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PREFACE
Preface
Although diagnosis always begins with a careful history and physical examination and
a physician is obligated to consider more than the diseased organ, testing of lung function
has become standard practice to confirm the diagnosis, evaluate the severity of
respiratory impairment, assess the therapy response and follow-up patients with various
cardio-respiratory disorders. Ventilation, diffusion, blood flow and control of breathing
are the major components of respiration and one or more of these functional components
can be affected by any disorder. Frequently, no single pulmonary function test yields all
the information in an individual patient and multiple tests have to be combined to allow
proper evaluation of the patient. The pulmonary function laboratory is therefore very
important in pulmonary medicine to provide accurate and timely results of lung function
tests.
The purpose of this issue of the European Respiratory Monograph is to provide up-todate information on the application and interpretation of different pulmonary function
tests in the work-up of patients suffering from cardio-respiratory diseases. In each
chapter of this issue, the contributors have attempted to relate theoretical considerations
of the different physiological tests to clinical application. New insights into the diagnostic
approach to patients with respiratory impairment form an integrated part of the different
chapters. This issue not only offers the reader a state-of-the-art approach to pulmonary
function testing, but also contributes significantly to a better understanding of the
pathophysiological processes underlying various diseases and contributing to the
morbidity of patients.
The guest editors of this issue, Henk Stam and Rik Gosselink, have done a great job in
the coordination and planning of this issue of the European Respiratory Monograph. The
authors of the different chapters have really tried to give the reader up-to-date
information about the different lung function tests. Therefore, I am convinced that the
knowledge and information provided in this issue of the European Respiratory
Monograph will contribute to the best possible evaluation and care for afflicted
individuals.
E.F.M. Wouters
Editor in Chief
Eur Respir Mon, 2005, 31, vii. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
vii
INTRODUCTION
H. Stam, R. Gosselink
The first indirectly described spirometer system consisted of a glass bottle without a
bottom, which was placed in a tub of water. The centre of gravity was so low that the
bottle did not capsize. The neck of the bottle was closed with a tap. The patient expired
via a tube, which led through to the underside of the bottle. Expiratory vital capacity
could be determined from the bottle’s displacement. There have been many changes since
these first descriptions of spirometry. Lung function research is a relatively young
science. Physicists have historically made an important contribution to the scientific
development of lung function analysis due to the importance of topics such as elasticity,
resistance, muscular strength and the work of breathing. These pioneers saw parallels
with models in electricity, with which they could calculate and predict lung function
results. However, the system of millions of alveoli and small airways are studied with a
relatively small number of indices, all measured at the mouth. In practice, simple models
appeared to give the most useful information. Nowadays, accurate measuring techniques
and the use of fast computers offer the pulmonologist lung function data that gives
specific information on, for example, airway resistance, ventilation equality, ventilationperfusion mismatch, diffusion characteristics of the blood-gas barrier, etc.
In this issue of the European Respiratory Monograph experts describe the state of the
art of a specific topic within the field of lung function. In each chapter, background,
technical possibilities and impossibilities, the importance in diagnosis and the
consequences for treatment are discussed. The measurement of lung function indices
in adults, as well as children, and the possibilities of measuring lung function in the
intensive care unit are described. The topics vary from simple office spirometry, as
performed by the general practitioner, to more sophisticated techniques, such as impulse
oscillometry performed in a lung function laboratory. Performing simple office
spirometry is not as simple as it seems. The spirometric indices are maximal
measurements and instruction is crucial. When equipment delivers a flow–volume
curve the appearance of the curve offers the general practitioner information on the
correctness of the measurement. Adults are relatively easy to instruct, but the instruction
of small children can be problematic. Measurements that do not require the cooperation
of the child are therefore preferable. An important development in paediatrics could be
the forced oscillation technique. In this method measurements are performed during
spontaneous breathing. With the help of superimposed pressure oscillations, information
on airway resistance is obtained. In spirometry the forced expiratory volume in one
second is an indirect measure of airway obstruction. In the Chapter 2 the measurement of
airway resistance using body plethysmography is described. The difference between total
lung capacity (TLC) obtained with the helium dilution technique and TLC obtained with
body plethysmography is a measure for trapped air. For a proper gas exchange alveolar
oxygen partial pressure needs to be high and carbon dioxide partial pressure low. The
ventilation process refreshes the alveolar gas breath-by-breath, while ventilation is
controlled by chemical and mechanical receptors. The arterial blood gas tensions provide
Eur Respir Mon, 2005, 31, viii–ix. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
viii
H. STAM, R. GOSSELINK
the simplest indicator of the adequacy of ventilatory control. Where there is little or no
mechanical abnormality, an elevation of the CO2 tension is an indication of inadequate
ventilation and impaired control mechanisms. The respiratory muscles play a crucial role
in the ventilation process. In Chapter 4 tests to evaluate the strength and endurance of
the respiratory muscles are described. The main function of the lung is the exchange of O2
and CO2 between the ambient air and the capillary blood. Diffusion characteristics of the
alveolo-capillary membrane and ventilation-perfusion mismatch play an important role
in gas exchange. These items are discussed extensively in Chapters 6 and 7. Finally,
exercise testing, where all the aforementioned systems are subjected to stress, is reviewed
in Chapter 8.
Unfortunately, a chapter dealing with reversibility and provocation tests in patients
with asthma to study hyperreactivity of the airways could not be included in this
Monograph. However, we are convinced that the most important issues concerning lung
function testing are reviewed.
ix
CHAPTER 1
Spirometry to detect and manage chronic
obstructive pulmonary disease and asthma
in the primary care setting
P.L. Enright*, M. Studnicka#, J. Zielinski}
*The University of Arizona, Tucson, Arizona, USA, #University Clinic of Pneumology, Paracelsus Private
Medical University, Salzburg, Austria, and }National Tuberculosis and Lung Diseases Research Institute,
Warsaw, Poland.
Correspondence: P. Enright, 4460 East Ina Road, Tucson, AZ 85718, USA.
Most people with chronic obstructive pulmonary disease (COPD) are unaware of the
smoldering airway inflammation present in their lungs, which places them at increased
risk for premature morbidity and mortality [1–3]. However, COPD is easily detected in
its preclinical phase using office spirometry; and successful smoking cessation prevents
further disease progression [4]. In the near future, other interventions may also be proven
to reduce the rapid decline in lung function experienced by patients with chronic airflow
limitation. When patients complain of intermittent cough, wheezing, chest tightness, and
shortness of breath, spirometry carried out when the symptoms remain current can often
detect the reversible airflow limitation characteristic of asthma. Spirometry also helps to
categorise the severity of asthma and confirms response to therapy [5]. Office spirometry
is defined as spirometry performed in the primary care (general practitioner) setting.
Office spirometry measures the forced expiratory volume in one second (FEV1)/vital
capacity (VC) ratio (or surrogates like FEV1/forced vital capacity (FVC) or FEV1/forced
expiratory volume in six seconds (FEV6)). This ratio is the most sensitive and specific test
for detecting airflow limitation. Spirometry also measures the per cent predicted FEV1,
which is the most widely accepted index of the severity of airway obstruction [6, 7].
General practitioners see the majority of adult smokers and patients with asthma, but
fewer than half use an office spirometer regularly [8, 9]. Barriers include the perceptions
that spirometers are expensive and difficult to use and maintain, that the test disturbs
patients and takes too much time to complete, that the reports are too difficult to
interpret, and that spirometry testing does not affect clinical outcomes.
Improvements in office spirometers
Recent improvements in spirometry hardware and software make it less expensive,
faster, and easier to obtain good quality spirometry test sessions, with automated
interpretations which aid clinical decision-making [10]. Pulmonary specialists and their
professional societies can use their knowledge and experience with pulmonary function
testing to help general practitioners to select a new office spirometer. Attempts to use
older spirometers often lead to frustration and abandonment by primary care
practitioners. Volume spirometers are too large, too expensive, risk cross-contamination,
and are difficult to maintain in the office setting. Older flow-sensing spirometers may
quickly become inaccurate as their sensors become clogged, and many lack quality
Eur Respir Mon, 2005, 31, 1–14. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
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P.L. ENRIGHT ET AL.
assurance software and modern reference equations [11]. Some new office spirometers are
as accurate as older volume spirometers [12].
Almost all spirometers that are sold now use an internal microprocessor or are
connected to a personal computer. See figure 1 for photographs of office spirometers.
The primary function of the computer is to measure the spirometry results for each
manoeuvre, calculate predicted values, and format a printed report. Office spirometry
software should also help the spirometry technologist to obtain better quality test
sessions [13, 14]. Each manoeuvre should be checked for acceptability and appropriate
error messages displayed (table 1). As additional manoeuvres are performed, the
repeatability of the FEV1 and FVC are determined, and a quality grade (A–F) computed
for the test session. The goal is to obtain an A or B grade by performing additional
acceptable FVC manoeuvres. An unbiased professional group will test the features of
Fig. 1. – Photos of several hand-held, battery operated, office spirometers.
Table 1. – Manoeuvre quality checks and test session quality grades
Acceptable manoeuvres:
Fast start (BEV v0.15 L)
Valid FEV6 (FET w6 s or FET 2–6 s with EOTV v0.04 L)
Test session quality grades
A = at least three acceptable manoeuvres,
with the largest two FEV1s matching within 0.1 L
and the largest two FEV6s matching within 0.1 L
B = at least two acceptable manoeuvres,
with FEV1s matching within 0.15 L
C = at least two acceptable manoeuvres,
with FEV1s matching within 0.2 L
D = only one acceptable manoeuvre (with no interpretation unless normal)
F = no acceptable manoeuvres (with no interpretation)
BEV: back extrapolated volume; FEV6: forced expiratory volume in six seconds; FET:
forced expiratory time; EOTV: end-of-test volume; FEV1: forced expiratory volume in one
second; There is no E grade specified for test quality (due to an academic tradition).
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OFFICE SPIROMETRY FOR GENERAL PRACTITIONERS
office spirometers, such as QC software, using a standardised checklist. The results will
be posted on the National Lung Health Education Program (NLHEP) website [15, 16] as
a guide to "consumers" who are planning the purchase of an office spirometer. A similar
service should be provided in Europe. See table 2 for a short list of desirable spirometer
features.
Six second manoeuvres
Office spirometry is faster and easier using six second manoeuvres. The six second
FVC (FEV6) is slightly smaller than the FVC (and the slow VC) when healthy persons
are tested, so reference equations for the FEV1/FEV6 and the FEV6 must be used [17, 18].
The FEV6 is more reproducible than the traditional FVC. The FEV1/FEV6 is just as
good as the traditional FEV1/FVC for diagnosing airflow limitation and for predicting
FEV1 decline in smokers [19, 20]. Short manoeuvres (without volume–time plateaus)
increase the risk of misclassification when traditional reference equations are used. The
use of six second manoeuvres reduces technologist and patient fatigue, and also
eliminates the risk of syncope when compared to prolonged FVC manoeuvres. However,
reference equations for the FEV6 are not yet widely available from European studies.
Until then, the traditional slow, inspiratory, or FVC may be used for the denominator in
the equation for the ratio FEV1/VC.
How to minimise misclassification
Unlike many medical tests during which the patient remains passive, spirometry
testing requires cooperation and an almost athletic breathing manoeuvre. With submaximal effort, the results are erroneous (false positive or false negative for disease or
change in severity). The misclassification rate is about 5% in most research and subspecialty settings, but the current authors experience is that misclassification has been
higher in primary care settings. The most common cause of error is inadequate
spirometry training and experience of the person performing the test [10, 11]. Instrument
inaccuracy or malfunction is much less frequently at fault.
The sources of variation of within-subject FEV1 measurements may be divided into
technical and biological components. The technical sources of error may be further
divided into those introduced by the instrument and those introduced by the interactions
between the technician and patient. Improvements in spirometry hardware and
manufacturing quality control, prompted by the development of clearly-defined
international standards, have reduced technical sources of variation due to the
instrumentation over the last decade. Checking volume accuracy using a 3.00 L syringe
Table 2. – Desirable features of new office spirometers
Only FEV1, FVC, and FEV1/FVC are reported
Automated manoeuvre quality checks
Test session quality grades (A–F)
Use of reference equations for six second manoeuvres
Disposable, reliable, inexpensive flow sensors
Flow–volume and volume–time curves are printed
Reports printed on plain paper
Automated interpretations
Rugged, battery power, 3 yr warranty
FEV1: forced expiratory volume in one second; FVC: forced vital capacity.
3
P.L. ENRIGHT ET AL.
filled with room air detects most sources of instrument drift and differences in the
accuracy of disposable flow sensors.
The primary source of variability is now the technician–subject interaction. Spirometry
tests, unlike electrocardiograms and venipuncture, require effort on the part of the
subject, prompted by directions from the technician. Each FVC manoeuvre requires
maximal effort during three phases of an "unnatural" breathing manoeuvre: 1) maximal
inhalation; 2) maximal exhalation for at least one second (for FEV1); and then 3)
continued exhalation for several seconds (for FVC). Submaximal inhalation effort
during the first phase reduces both the FEV1 and the FVC. A submaximal exhalation
blast during the second phase affects the FEV1; and an incomplete (short) exhalation
during the final phase will reduce the measured FVC. Any (and sometimes all) of these
three phases of the manoeuvre can go wrong, usually because of suboptimal
communication between the technician and the subject, but sometimes because of
fatigue, lack of interest, or poor mental function. See figure 2 for examples of poor
quality spirometry manoeuvres.
The current European Respiratory Society (ERS) and American Thoracic Society
(ATS) goals for spirometry quality (three acceptable manoeuvres, the best two of which
are reproducible) [21, 22] are not unrealistic, at least in the hospital-based pulmonary
function testing (PFT) laboratory and research settings. Ninety-five per cent of 18,000
tests of adult patients, performed by 16 technicians in a very large clinical PFT lab, met
ATS standards [23]; and 95% of 4,000 tests of elementary and high school students (aged
9–18 yrs) performed by 12 different technicians in a research study, also met ATS
standards [24]. Tests of patients with asthma enrolled in six large multicentre asthma
research studies at 232 sites also met ATS goals [25]. Even nine out of 10 tests in elderly
people at their first research study visit could meet ATS standards [26].
A recent study in The Netherlands compared the spirometry results carried out by 388
patients with mild-to-severe COPD first tested in four hospital-based PFT laboratories
with repeat studies carried out in 61 general practice outpatient clinics [27]. The same
12
10
C
Flow L·s-1
8
D
6
4
A
2
0
0
1
B
2
3
Volume L
4
5
6
Fig. 2. – Examples of the patterns of common spirometry errors causing misclassification. A: a hesitating start
(––); B: a submaximal blast (– ? –); C: large coughs during the first second (?????); D: quit too soon (-----).
4
OFFICE SPIROMETRY FOR GENERAL PRACTITIONERS
Table 3. – Factors to consider during interpretation of
spirometry results to minimise misclassification
The pre-test probability of disease
The patient’s risk factors (age, sex, symptoms, etc.)
The quality of the test session (often graded A–F)
The distance from the LLN (% predicted)
The consequences of a falsely-positive interpretation
The consequences of a falsely-negative interpretation
LLN: lower limit of the normal range.
model of office spirometer was used at all locations. The mean FEV1 and FVC results
were nearly identical when repeated, but the individual results differed by up to 0.4 L for
FEV1 (5–95th percentile confidence intervals) and up to 0.8 L for FVC. Furthermore, in
both settings, 18% of the tests did not meet ATS standards, and the investigators
concluded that perhaps their "gold standard" (testing done in PFT labs) was actually a
"gilded standard". The office nurses (with a mean of 11 yrs of experience) were centrally
trained using two 2.5 h courses, 1 month apart, but there was no over-reading and
reporting system. The office spirometers had calibration checks done every 3 months.
Spirometry training materials are available on the Internet [28] and CD-ROM [29]. The
current authors recommend that professional societies develop office spirometry
certification programmes for nurses and technologists, which are based on practical
knowledge and demonstrated performance of good quality spirometry tests.
The accuracy of a test for screening or case-finding is measured in terms of two indices:
sensitivity and specificity. A test with poor sensitivity will miss cases, producing falsely
negative results, while a test with poor specificity will result in healthy persons being told
that they have the disease (a falsely positive result). The sum of the false negative rate and
the false positive rate is the overall misclassification rate. Five per cent is usually
considered an acceptable misclassification rate for most medical tests; thus one in twenty
patients will get an inaccurate interpretation of the test results. See table 3 for factors to
consider during the interpretation of spirometry results to minimise the risk of
misclassification. See table 4 for a list of methods to minimise the misclassification rate.
A recent recommendation suggests that 70% is used as the lower limit of the normal
range (LLN) for the ratio FEV1/FVC [30]. However, use of a fixed LLN will increase the
Table 4. – Tips for interpreting office spirometry results
1. Poor quality test sessions often cause diagnostic misclassifications.
2. First look at the pattern of the curves, then the numbers to confirm your impression.
3. A bowl or rat’s tail shaped flow–volume curve suggests airways obstruction. A low ratio confirms airway
obstruction.
4. A normal flow–volume curve looks like a sail, rising rapidly to a peak, then descending at about a 45 degree
angle.
5. If the volume–time curve stops before 6 s and doesn’t reach a flat plateau, the FVC (and FEV6) are
underestimated.
6. A low FVC with a normal ratio suggests restriction without obstruction. Restriction may be verified by
measurement of total lung capacity.
7. In a patient with respiratory symptoms, airway obstruction with an FEV1 which increases by w12% (and w0.2 L)
suggests asthma.
8. In a patient with intermittent respiratory symptoms, the lack of airway obstruction, or the lack of a bronchodilator
response do not rule out asthma.
9. Airway obstruction in an adult smoker is usually (but not always) due to COPD.
10. After spirometry, if you remain uncertain of the diagnosis, consider a diffusing capacity test (for emphysema or
interstitial lung disease) or a methacholine challenge test (for asthma).
FVC: forced vital capacity; FEV6: forced expiratory volume in six seconds; FEV1: forced expiratory volume in one
second; COPD: chronic obstructive pulmonary disease. Adapted from [68].
5
P.L. ENRIGHT ET AL.
misclassification rate when detecting airflow limitation. Instead, the LLN should be age
and sex-specific. All published population-based studies of spirometry show that the
ratio decreases with age in the healthy subset of the population, suggesting that aging
alone causes slightly progressive airflow limitation (fig. 3). While 70% is about right for a
50-yr-old male, the 5th percentile LLN for a 20-yr-old is about 75%, and for an 80-yr-old
65%. The use of a fixed 70% threshold causes considerable misclassification when applied
to either young adults (where the false-negative rate becomes high) or elderly adults
(where the false-positive rate becomes high) [31].
Accept uncertainty
Clinicians much prefer to view test results as black-or-white, abnormal or normal, but
such a stubborn stance increases the misclassification rate. Results that are near the
rather arbitrary threshold (the LLN) should instead be interpreted with uncertainty
(fig. 4). For instance, if the LLN for the FEV1/FVC is 73% and the patient’s ratio is 72, it
should not be stated with confidence that a smoking patient has airflow limitation and
COPD. On the other hand, if the patient’s ratio is 55% (and the patient’s FEV1 is 60%
pred) even if the quality of the spirometry test was suboptimal, one can state with
confidence that the patient has COPD. Changes in the FEV1 due to therapeutic
interventions which are near the threshold of clinical significance should also be
considered "borderline" (of uncertain significance).
The 2003 Global Initiative for Obstructive Lung Disease (GOLD) document correctly
emphasises that "maximal patient effort in performing the test is required to avoid errors
in diagnosis and management" and that "the supervisor of the test needs training in its
effective performance" [30]. The National Lung Health Education Program (NLHEP)
document goes much further by requiring that office spirometers incorporate software
that automatically checks manoeuvre acceptability and then checks for repeatable FEV1s
and FVCs before the test session is considered complete [10]. It also recommends that
manufacturers take an active role to enable office staff to learn how to use their
FEV1/VC × 100%
90
85
M
80
F
75
M
70
F
Mean
Lower limit
65
60
55
30
40
50
Age
60
70
80
Fig. 3. – The forced expiratory volume in one second (FEV1)/forced vital capacity (FVC) decreases with age
(figure shows normal (predicted) FEV1/VC from the third National Health and Nutrition Examination Survey
(NHANES III)). Using a fixed ratio (like 70%) to determine airway obstruction will cause misclassification in
young people and the elderly. M: male; F: female.
6
OFFICE SPIROMETRY FOR GENERAL PRACTITIONERS
High confidence
Low confidence
FEV1/FVC:
50%
Abnormal
“Black”
70%
LLN
“Grey”
90%
Normal
“White”
Fig. 4. – Confidence in spirometry interpretation should be low when the forced expiratory volume in one
second (FEV1)/forced vital capacity (FVC) (or the vital capacity) are near the lower limit of the normal range
(LLN).
spirometer by providing easy-to-understand educational materials, such as audio-visual
aids.
Spirometry for finding cases of chronic obstructive pulmonary
disease
The authors recommend office spirometry for COPD case-finding when the following
five factors are true: 1) an adult patient is seen in a healthcare setting; 2) the patient is a
current or former smoker, especially those with any respiratory symptom (chronic cough,
phlegm, wheezing, or dyspnoea on exertion); 3) good quality spirometry is carried out; 4)
the result is interpreted correctly; and 5) the patient is referred to an effective local
smoking cessation programme [32, 33]. Undiagnosed airflow limitation (airway
obstruction) is common in the general population, is associated with impaired health
and functional status [3, 34–36], and is an independent predictor of morbidity and
mortality [37]. Airflow limitation due to smoking is unusual at agev40 yrs. The presence
of any respiratory symptom doubles the risk of airflow limitation [38]. Simple
measurement of peak flow cannot substitute for spirometry, either for detecting
airway obstruction or determining its severity [39, 40].
The most common causes of airway obstruction are asthma and COPD [1, 31]. While
almost all hospitals have a PFT laboratory, and almost all pulmonary specialists own a
spirometer, the current authors estimate that v50% of general practitioners actively use
spirometry in their practice. In the USA, the NLHEP promotes the appropriate use of
spirometry by general practitioners for the detection of COPD in smoking adults [10].
However, "screening" for COPD remains controversial [42–45] since it has not yet been
proven that the staff in the offices of general practitioners can attain the same low
misclassification rate as experienced and certified pulmonary function technologists who
perform spirometry in a laboratory setting [39, 40].
Relatively healthy workers between the ages of 40–50 yrs are less likely than older
adults to visit a physician. Spirometry performed in the workplace setting may detect
COPD in this age group more frequently than waiting for them to be seen by a physician.
Another alternative is to invite smokers in this age group to call for an appointment for
spirometry testing at a convenient clinic or pulmonary function laboratory. In one such
study, COPD was detected in about one-quarter of those who responded [46].
There is a difference between using medical tests for screening versus case-finding. An
example of screening is a booth at a city festival or sporting event which offers to perform
7
P.L. ENRIGHT ET AL.
spirometry for anyone who is interested [47]. An example of case-finding is a physician
who performs spirometry during an office visit for a patient with an unrelated disorder,
such as hypertension. For example, the review of systems of a 50 yr-old female patient
may disclose current smoking and a chronic morning cough, a combination of COPD
risk factors that provides a clinical indication for spirometry testing. The physician then
discusses the spirometry results with her and refers her to a local smoking-cessation
programme. Spirometry for COPD case-finding in adult cigarette smokers fulfills all of
the standard criteria for application of medical test for screening [48]; however, the
evidence for two of these criteria remains weak. While spirometry is indeed accurate in
the PFT lab setting (has a low misclassification rate), this may not be true in some
outpatient settings [49]. It has been shown that adding spirometry to an optimal smoking
cessation programme statistically significantly increases the subsequent 12 month
smoking cessation rate [50–52]. Although the slightly higher rate may not be noticed by a
single general practitioner [53], even a 2% improvement in smoking-cessation rates (for
example, from 10% to 12%) would result in a very large absolute number of lives saved
every year in a single country [54]. Of course, primary prevention of COPD, by
prompting children to avoid becoming addicted to cigarette smoking and reducing
workplace air pollutants, is even more important than secondary prevention efforts such
as case-finding.
Potential adverse effects of screening for chronic obstructive
pulmonary disease
There are tangible and intangible costs of any medical test. Adverse effects may occur
due to: 1) the procedure itself; 2) the investigation of abnormal results; or 3) the
treatment of detected abnormalities or diseases [48]. The economic cost of spirometry
includes the cost of the instrument and the cost of personnel time (both training and
testing). Office spirometers currently cost about J1,000 and about J10 of time per test is
spent for testing (including initial training time) and disposable supplies. The authors
estimate that accurate office spirometers will soon costvJ500. There are no adverse sideeffects from the test itself, other than occasional minor discomfort that lasts for a few
minutes.
Investigation and confirmation of abnormal spirometry results consumes both time
and money, and may result in psychological and social harm to a few. The cost of
diagnostic spirometry to confirm airflow obstruction, when performed in a hospitalbased PFT lab is substantial. The estimated travel time, waiting time, and testing time
spent by the patient ranges from 1 h to 3 h. The possible psychological impact of being
labeled as "ill" by self and others related to false positive or even true positive test could
lead to alterations in lifestyle, work, and seeking medical attention. Another important
potential adverse effect is the unmeasured risk of reinforcing the smoking habit in some
of the four out of five adult smokers who are told that they have normal spirometry.
However, physicians should counteract this possibility by taking the opportunity to tell
the patient that although spirometry was normal, their risk remains high of dying from a
heart attack, lung cancer, and other smoking-related diseases; therefore, smoking
cessation remains very important.
The risk of an adverse effect caused by smoking cessation is very small, and the side
effects of nicotine replacement therapy and bupropion are minor. Successful smoking
cessation leads to an average increase in body weight [55], but the slight increase in
medical risk from minor weight gain is far exceeded by the benefits due to reduced
morbidity and mortality. On the other hand, if long-acting bronchodilators or
8
OFFICE SPIROMETRY FOR GENERAL PRACTITIONERS
corticosteroids are inappropriately prescribed, the cumulative cost is high and the
potential side-effects can be very serious in elderly patients [56–58].
The chronic obstructive pulmonary disease action plan
Early intervention following early identification of lung function abnormalities can
lead to improved smoking cessation, workplace or home environmental changes, and
increased awareness and attention to cancer, cardiac and other nonpulmonary health
issues associated with abnormal lung function. Early identification of airway obstruction
in relatively asymptomatic patients may provide "teachable moments" when the patient
has an increased awareness and response to medical education and intervention [59]. The
patient is more likely to consider smoking cessation again.
Once an abnormality has been detected, an action plan must follow. Repeat
spirometry should be performed to confirm abnormal office spirometry prior to initiating
an expensive work-up, or when considering interventions with negative economic
consequences (such as expensive medications or a recommendation to change jobs).
When airway obstruction is identified in a smoker, the primary intervention is smoking
cessation, since it is currently the only intervention that has been demonstrated to halt
rapid decline in lung function, and thereby reduce the risk of disabling COPD [58]. In
smokers with airway obstruction but without dyspnoea on exertion, smoking cessation is
the only intervention with proven value. Referral to a subspecialist for further diagnostic
testing should be considered in some cases. In the event that a patient with airway
obstruction continues to smoke cigarettes, renewed/increased effort to assist with
smoking cessation is essential [60].
Spirometry for confirming and managing asthma
Asthma is very common at all ages, and the symptoms are often overlooked or
mistakenly attributed to other problems. Asthma is a disease of airway inflammation,
airway hyperresponsiveness, and intermittent airway obstruction. Office spirometry can
easily detect airway obstruction in a patient with asthma who presents to a primary care
practitioner with respiratory symptoms such as chronic cough, chest tightness, or
wheezing [61]. The FEV1 (compared to the predicted value) may then be used to help
classify the severity of asthma [7].
Since the airway obstruction of asthma is intermittent, a normal FEV1/FVC during a
single visit does not rule out asthma: referral for an inhalation challenge test should then
be considered to confirm or rule out asthma. If baseline spirometry shows airway
obstruction, it should then be repeated 10–15 min after inhalation of salbutamol, to
detect bronchodilator responsiveness. An increase of at least 12% and 0.2 L in the FEV1
(baseline or predicted) helps to confirm asthma [6] and predicts a good response to
asthma therapy; however, the lack of acute improvement with bronchodilator inhalation
does not rule out asthma. A clinical trial of asthma controller medication (4–8 weeks)
should be considered, with repeat spirometry at the follow-up exam.
Office spirometry for measurement of treatment responses
An important goal of asthma management is to keep lung function close to the
patient’s personal best value (the green zone of good control). Asthma controller
9
P.L. ENRIGHT ET AL.
medications should be stepped up to reach this goal and then stepped down, while
monitoring to ensure that the patient remains in the green zone. No single asthma
controller medication works well for all patients with asthma. Some patients may not
respond to inhaled corticosteroids; some do not respond to leukotriene antagonists;
while others do not respond well to long-acting bronchodilators [62]. This means that
objective evidence for the effectiveness of these expensive medications (some with serious
side-effects) should be sought during follow-up visits. Spirometry should be used to
supplement the results from an asthma diary and responses to questions about the
frequency of nocturnal awakenings and need for rescue medication. An improvement of
w15% in FEV1 from one visit to the next is clinically significant. Changes in peak flow are
less sensitive and less specific for detecting change in lung function when compared to
following changes in the FEV1 [63].
Spirometry is also useful for determining the response of bronchodilator therapy given
for relief of dyspnoea in patients with COPD. Improvement in the FEV1 remains a
primary outcome measure for most COPD clinical trials [64]. An improvement of more
than 0.3 L in FEV1 from one visit to the next is outside of the noise of measurement [65]
and clinically significant in patients with mild-to-moderate COPD (an FEV1 above 50%
pred). However, following changes in the FEV1 is probably not helpful in individual
patients with COPD whose FEV1 is severely reduced (below 1 L).
Examples of spirometry testing programmes
A national programme of early diagnosis and prevention of COPD in Poland has been
reported [66]. It started in 2001 in 12 cities, where over 11,000 ever-smokers were tested in
pulmonary outpatient clinics. About one-fourth of those tested had airflow limitation
(10% mild, 10% moderate, 5% severe). They were all given advice to stop smoking by a
physician. About 9% had the nonspecific pattern of a low FVC without airway
obstruction. Two-thirds of the participants returned for a follow-up visit about 12
months later [52]. Half of those who returned had airflow limitation during their baseline
exam. The biochemically verified 12 month smoking-cessation rates showed that those
with moderate-to-severe airflow limitation were twice as likely to have quit when
compared to those without airway obstruction (17% versus 8.4% quit rates). The
independent predictors of success were a late start of smoking, older age, fewer packyears, and a lower FEV1. There was no sex difference in quit rates.
Two programmes of asthma and chronic obstructive pulmonary disease screening
were completed in The Netherlands [53, 67]. From two semi-rural general practice offices,
spirometry testing was carried out for 651 adult current smokers. According to American
Thoracic Society criteria, 85% had acceptable test session quality, and of those, 18% had
an abnormally low forced expiratory volume in one second. Patients reporting a chronic
cough were about twice as likely as the other smokers to have abnormal spirometry; and
nearly half of the smokersw60 yrs had abnormal spirometry. The authors estimated that
in each practice, when one adult smoker was tested every day, one case of chronic
obstructive pulmonary disease was found per week.
Summary
Office spirometry in the primary care setting can be most helpful for the detection
(case finding) and management of asthma and chronic obstructive pulmonary disease
10
OFFICE SPIROMETRY FOR GENERAL PRACTITIONERS
(COPD). The severity of asthma is underestimated by history and physical
examination alone in some patients. Only spirometry has been shown to detect
COPD in its early stages. The cost and side-effects of medications for asthma and
COPD drives the need for objective measurement of their response, by measuring the
forced expiratory volume in one second during follow-up visits. The value of
population-based screening for these diseases needs further evidence. The new
generation of office spirometers are less expensive, include quality checks, and make
spirometry easier using six second manoeuvres. However, enthusiastic coaching for
correct breathing manoeuvres remains important to reduce the risk of misclassification, which is substantial in the primary care setting.
Keywords: Airflow limitation, airway obstruction, asthma, chronic obstructive
pulmonary disease, smoking, spirometry.
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14
CHAPTER 2
Whole-body plethysmography
M.D. Goldman*, H.J. Smith#, W.T. Ulmer}
*David Geffen School of Medicine, University of California, Los Angeles, USA. #Research in Respiratory
Diagnostics, Berlin, Germany. }Research in Lung Function, Bochum, Germany.
Correspondence: M.D. Goldman, David Geffen School of Medicine, University of California, Los Angeles,
USA.
The word plethysmograph is derived from the Greek plethusmos (enlargement), and is
related closely to plethus (fullness) and plethora (fullness). Indeed, the fundamental
function of a whole-body plethysmograph is the measurement of intrathoracic gas
volume (TGV) and volume change. Whole-body plethysmographs have been used to
measure changes in lung volume over a range of volumes, from the scale of millilitres to
litres. Early reports of whole-body plethysmography to determine thoracic gas volume
(TGV) [1] and airway resistance (Raw) [2] measured volume changes of the order of
millilitres, in terms of associated changes in plethysmograph and alveolar pressures (Pa),
using the constant-volume variable-pressure plethysmograph. Changes in lung volume
during compression and decompression of thoracic gas were measured while the subject
breathed entirely within the plethysmograph.
An alternative volume-displacement whole-body plethysmograph measured volume
changes of the thorax directly, including both changes in volume of gas flowing into and
out of the lung and simultaneous changes in compression and decompression of thoracic
gas [3]. In contrast to the constant-volume plethysmograph of DuBois et al. [1], subjects
breathed in and out across the wall of the volume-displacement plethysmograph
developed by Mead [3]. The volume-displacement plethysmograph provided more ready
assessment of changes in TGV during extended manoeuvres such as the vital capacity
(VC). During such forced manoeuvres, lung volume changes due to compression of
thoracic gas were measured, in addition to those associated with gas flow out of the lung.
Subsequent technological developments permitted a combination of the two
approaches by using a pressure-compensated volume-displacement or integrated-flow,
plethysmograph, described by the groups of Mead and van de Woestijne [4–7], and
reviewed by Peslin [8] and Coates et al. [9]. In the combination plethysmograph, the
subjects breathe either across the wall of the plethysmograph to the outside to measure
total thoracic displacements or within the plethysmograph to measure compression
volumes only, excluding air flow into or out of the lung. In this combination
plethysmograph, both pressure change in the plethysmograph and the volume displaced
through the plethysmograph wall are combined to provide a measure of the volume
displacements of the thorax. This approach provides the advantageous frequency
response of the pressure plethysmograph with the ability to measure volume
displacements over a very wide range of volumes. This approach is now commonly
referred to as a "transmural" plethysmograph.
Current technological improvements in whole-body plethysmography provide
measurable variables that are less dependent on patient cooperation than in initial
implementations [1, 2]. Recent advances in the understanding of chronic obstructive
pulmonary disease (COPD) have led to renewed interest in the evaluation of compression
Eur Respir Mon, 2005, 31, 15–43. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
15
M.D. GOLDMAN ET AL.
of TGV as an aid to better understanding of dynamic events during the respiratory cycle.
Therefore, applications of plethysmographic techniques that include thoracic volume
displacements are reviewed as well in this Chapter. However, this Chapter focuses
primarily on the use of the variable-pressure constant-volume plethysmograph, as this
instrument has been in use most commonly in clinical pulmonary function testing. The
clinical measurements of Raw and functional residual capacity (FRC) determined by
whole-body plethysmography (FRCpleth) are most extensively discussed herein.
Principles of whole-body plethysmography
The whole-body plethysmograph consists of a rigid chamber, of comparable size and
shape to an enclosed telephone booth, in which the subject sits while breathing through a
pneumotachograph. Pressure transducers of different sensitivity are arranged to measure
the pressure across the pneumotachograph (flow), the pressure difference across the wall
of the plethysmograph and pressure at the airway opening. The fundamental principle of
the variable-pressure plethysmograph is that changes in PA may be inferred from
changes in plethysmograph pressure. This is achieved by the process described
immediately below.
A shutter mechanism is positioned close to the mouth in the plethysmograph. This
shutter may be closed to provide transient airway occlusion. Voluntary respiratory
efforts are performed against the closed shutter, during which the change in PA (DPA), is
estimated by recording the change in mouth pressure (DPm). Pm (PA) is plotted against
simultaneous plethysmographic pressure changes during respiratory efforts against a
closed shutter to measure absolute TGV. The same relationship between alveolar and
plethysmographic pressure measured during respiratory efforts against a closed shutter is
then extended to dynamic events during free breathing to measure Raw, wherein airflow is
related to PA.
Types of plethysmograph
Three different types of whole-body plethysmograph may be used to measure changes in
thoracic volume. These depend on whether the aim is to measure the large volume changes
associated with respiratory manoeuvres, such as the VC, or just those which accompany
compression and decompression of the gas in the lungs exclusive of changes in volume due
to gas flow in and out of the lungs. Suitable changes in transducer sensitivities and
mechanical arrangement are incorporated into the different types of plethysmograph.
The constant-volume or variable-pressure plethysmograph is used to measure small
volume changes due to compression and decompression of gas within the lungs.
The constant-pressure or volume-displacement plethysmograph is used to measure
large changes in lung volume associated with gas flow into and out of the lungs.
The pressure-corrected variable-volume plethysmograph combines the advantages of
both the plethysmographs described above. The sensitivity and rapid frequency response
of variable-pressure constant-volume plethysmography is provided along with the ability
to measure large slow volume changes in the lungs during breathing.
Variable-pressure plethysmograph. The advantage of the variable pressure plethysmograph is simplicity of hardware components and accuracy of the measurement. The
small changes in plethysmographic pressure associated with compression/decompression
of TGV are recorded using a very sensitive pressure transducer, as shown schematically in
figure 1.
16
WHOLE-BODY PLETHYSMOGRAPHY
Fig. 1. – Schematic representation of a variable-pressure constant-volume plethysmograph, illustrating the
controlled mechanical leak to room air and reference chamber. The subject breathes through a pneumotachygraph, entirely within the chamber. Recording of the volume displacements of the thorax is limited to those
related to compression and decompression of thoracic gas. Calibration of plethysmographic pressure is done via
a motorised syringe inserting and withdrawing 30–50 mL of air into the plethysmograph chamber at a frequency
of approximately 1 Hz. PA: alveolar pressure; Pm: mouth pressure; VL: lung volume; Raw: airway resistance;
TGV: thoracic gas volume; DV: change in volume; DP: change in pressure. See text for discussion.
The plethysmograph is open to the atmosphere via a small leak with a mechanical time
constant of 5–25 s, with most current instruments using a value v10 s. This controlled
leak minimises slowly occurring pressure changes that are not related to respiratory
manoeuvres, such as thermal drift (heating) caused by the presence of a subject breathing
within the chamber.
The large volume of the plethysmograph chamber (600–1,000 L) undergoes very small
pressure changes during compression and decompression of TGV. Accordingly, the
plethysmographic pressure transducer must be very sensitive and stable. It is stabilised
against changes in room air pressure during such events as opening or closing of a door
by connection of the other side of the plethysmographic pressure transducer to a
reference chamber with comparable time-constant to that of the plethysmograph.
In practice, the plethysmographic pressure transducer is calibrated in terms of changes
in TGV. This is done by quickly introducing and withdrawing 30–50 mL air into the
plethysmograph using a motor-driven syringe, to simulate the changes in TGV that occur
during decompression and compression of thoracic gas. After such calibration, the
measured changes in plethysmographic gas pressure reflect the change in TGV due to
compression and decompression of thoracic gas. Changes in calibrated plethysmographic
gas pressure are recorded in terms of volume change (DV), and known as shift volume.
Shift volume is the change in TGV due only to compression or decompression, exclusive
of changes due to airflow into and out of the lung, both during occluded respiratory
efforts and during breathing within the plethysmograph. Since calibration of the
plethysmograph is normally carried out without a subject in the plethysmograph, this
calibration must be corrected for the subject’s body volume. Therefore, the body weight
of the subject is entered prior to any testing of the subject and is used in the calculation of
the final calibration coefficient.
Volume-displacement plethysmograph. A volume-displacement plethysmograph
measures volume changes of the thorax directly. Subjects breathe in and out across
17
M.D. GOLDMAN ET AL.
the wall of the plethysmograph to room air. The increase in lung volume that occurs
during inspiration includes the volume of gas inspired plus the additional volume
associated with decompression of TGV resulting from the fall in intrathoracic pressure
necessary to provide a gradient for inspiratory airflow.
The advantage of the volume-displacement plethysmograph is the ability to measure
respiratory manoeuvres such as the slow or forced vital capacity (FVC). Integrated
airflow at the mouth can be compared to thoracic volume displacements during forced
expiration to provide more physiological information in subjects with hyperinflation or
airway obstruction.
Measurement of total thoracic volume displacement is useful, but the original plethysmograph described by Mead [3] required a very sensitive and critically damped direct-reading
spirometer, which was technically very demanding to build. Therefore, this construction has
been supplanted by the pressure-corrected integrated-flow plethysmograph [6–9].
Flow plethysmograph. Comroe et al. [10] pointed out that use of a spirometer connected
to a volume-displacement plethysmograph chamber made it difficult to obtain adequate
speed of response. Frequency response was improved by adding a signal proportional to
plethysmograph pressure to the volume-displacement signal and, subsequently,
substituting a pneumotachograph in the wall of the plethysmograph for the
spirometer bell [4–6]. Such a pressure-corrected plethysmograph which integrates flow
through the plethysmograph wall permits accurate measurement of changes in TGV
during forced expiration manoeuvres. Loss of sensitivity with small thoracic volume
displacements and possible zero-flow integrator drift are limitations of this approach
when measuring TGV; but occlusion of the pneumotachograph in the wall of the
plethysmograph converts the flow plethysmograph back into a variable-pressure
plethysmograph, allowing more sensitive measurements of TGV. The flow plethysmograph is shown schematically in figure 2.
Some pressure change in plethysmograph air is required to cause movement of air in
Fig. 2. – Schematic representation of a pressure-corrected integrated-flow plethysmograph, illustrating the path of
the subject’s breathing through the wall of the plethysmograph to room air. Recording of the volume
displacements of the thorax includes those due to airflow into and out from the lung, as well as those due to
compression and decompression. Volume displacements of the thorax drive plethysmograph air through a
pneumotachygraph in the wall of the plethysmograph, and are recorded by integration of the flow through the
plethysmograph wall. PA: alveolar pressure; Pm: mouth pressure; VL: lung volume; Raw: airway resistance; TGV:
thoracic gas volume; DV: change in volume; DP: change in pressure. See text for further discussion.
18
WHOLE-BODY PLETHYSMOGRAPHY
and out of the plethysmograph chamber. This pressure change occurs in the large volume
of compressible gas within the plethysmograph chamber. Thus, part of the volume
displacement is temporarily "lost" in compression or decompression of plethysmographic
air and does not reach its equilibrium value until plethysmographic air pressure has
returned to atmospheric, as noted by Mead [3]. This volume displacement is "found"
within the pressure change of plethysmographic air itself. Thus, as the subject breathes
room air through a tube across the wall of the plethysmograph, changes in TGV expand
or compress plethysmographic air, and simultaneously displace some air in or out of the
plethysmograph across the flow meter in its wall. The volume displaced by compression
or decompression of plethysmograpic air is recovered by adding an electrical signal
proportional to plethysmographic pressure to the measured volume displaced across the
plethysmograph wall in the "pressure-corrected" body plethysmograph [4–7]. Because
this volume displacement is most commonly recorded by integrating flow through a flow
meter in the plethysmograph wall, this type of plethysmograph is often described as a
pressure-corrected integrated-flow plethysmograph. It should be emphasised that in this
use, the integrated flow is the flow in and out across the wall of the plethysmograph
chamber, in contrast to the integrated airflow in and out of the mouth described for the
pressure plethysmograph [2].
The measurement of a rapid volume change, such as that encountered during a brief
cough or the initial part of a forced expiratory VC utilises the "pressure-correction"
shown schematically in figure 3 which is modified from Leith and Mead [7]. Figure 3
shows schematically the initial rapid decrease in TGV during a forced expiration. The
trace labelled a) represents an idealised trace of the true volume change for the initial
rapid decrease in lung volume, shown as the onset of a square wave. At the onset of this
abrupt decrease in TGV, plethysmographic pressure falls rapidly with compression of
thoracic lung volume and expansion of plethysmographic air by expiratory muscle effort,
then exponentially returns to its initial value, after the volume event is complete (e.g. a
brief cough). This pressure change is shown in trace b). The signal from a linear
flowmeter in the wall of the plethysmograph is identical in shape to trace b). Integration
of this flow signal is shown in trace c). Integrated flow across the plethysmograph wall
Time
a) True
volume change
b)
c)
Volume
Ppleth
V´pleth
òV´
pleth
Ppleth
d)
ò
òV´
Ppleth + V´pleth
pleth
Fig. 3. – Schematic representation of the basis of "pressure-correction" to account for phase lag between volume
displacements of the thorax and those of plethysmograph air through the pneumotachygraph in the
plethysmograph wall. Ppleth: pressure in the plethysmograph chamber; V9pleth: flow across wall of chamber. See
text for discussion.
19
M.D. GOLDMAN ET AL.
eventually reaches the same level as the true decrease in TGV, but the volume change
recorded by integrated flow across the plethysmograph wall is slower because of the
temporary "loss" of volume during the initial decrease in plethysmographic air pressure.
To recover this contribution, an electrical signal proportional to trace b) is added to the
integrated plethysmograph flow in trace c). The sum of these contributions recovers the
initial true volume event represented by the solid trace d).
Pressure-corrected integrated-flow plethysmographs provide sensitive recordings of
pressure and volume events over a wide range of volume displacements. They permit
accurate recording of maximal expiratory flow-volume curves in addition to measurement of TGV, specific airway resistance (sRaw) and Raw with the same instrument. Thus,
this approach provides the advantageous frequency response and sensitivity of the
variable-pressure plethysmograph with additional lung volume displacement recordings
over a wide range of volumes, and has been used for measures of true TGV change
(including that due to compression of thoracic gas) during either tidal breathing [11] or
measurement of the FVC [12]. This combined approach is now commonly referred to as a
"transmural" plethysmograph. It permits evaluation of the differences between thoracic
gas compression and airway closure, the so-called "trapped gas" [13].
Measurement notes
Applications of whole-body plethysmography include physiological evaluation of
respiratory mechanical limitations and diagnostic clinical testing. Special applications
include paediatric and infant diagnostic testing, which have been extensively discussed by
others [9, 14]. The present chapter is restricted to plethysmographic measurements in
larger children and adults. While spirometry is the most commonly used pulmonary
functional diagnostic test, body plethysmography provides essential additional
diagnostic information [15, 16], and usefully includes measurement of both slow and
forced vital capacities done in the plethysmograph.
After the subject has entered the plethysmograph, the door is closed with an airtight
seal. Approximately 2 min are required for plethysmograph cabin pressure to equilibrate
while air in the cabin is warmed and humidified by the subject breathing at rest. During
this initial period, the plethysmograph cabin is vented periodically to room air via a
solenoid-operated valve. After about 2 min, pressure drift with the valve closed is much
decreased and does not interfere with the measurement of sRaw. At this time the subject is
asked to close his/her lips tightly around the mouthpiece and breathe normally through
the pneumotachygraph. The patient sits erect with head and neck in a neutral posture. A
nose-clip is applied to close the nares. The subject is allowed to adapt to the measurement
conditions and breathe regularly through the flowmeter for about 30 s before testing is
initiated.
A complete whole-body plethysmography, measurement is commonly divided into
three standardised measuring sequences whose order may be defined by diagnostic
requirements. sRaw is usually measured first, followed by measurement of TGV and
concluding with measurement of the entire range of lung volumes, both slow and forced
spirometry. Individual measurement phases can be skipped or repeated, depending on
the diagnostic information required and/or the patient’s ability to cooperate. The initial
report of whole-body plethysmography first described its application to measure TGV [1]
and the description that follows below begins with determination of TGV. The
measurement of lung volumes by plethysmography is extensively reviewed by Coates
et al. [9] along with detailed discussions of physiological assumptions and technical
demands of measuring instruments.
20
WHOLE-BODY PLETHYSMOGRAPHY
Determination of thoracic gas volume and functional residual capacity
As befits the etymology of body plethysmography, its primary use to measure TGV
is considered first. It is noted that while this order corresponds didactically to the
historical development of plethysmography and previous literature, in practice, current
computer-assisted plethysmographic techniques commonly measure specific resistance
first.
Measurement of TGV is done in the variable-pressure constant-volume plethysmograph by making use of Boyle-Mariotte’s law which relates pressure and volume changes
to each other under isothermal conditions. Thus, during compression of thoracic gas, its
pressure rises and, at constant temperature, the product of pressure and volume remains
constant. In the plethysmograph, voluntary respiratory efforts are used to produce
changes in alveolar dry gas pressure, DPA, which are associated with reciprocal changes
in TGV, DV. Alveolar dry gas pressure (PA) itself is the difference between ambient
barometric pressure (Pbar) and saturated water vapour pressure at body temperature
(PH2O,sat), when the glottis is open with no airflow. A shutter mechanism positioned close
to the mouth provides for transient controlled airway occlusion, which is utilised in
making voluntary respiratory efforts to determine the relationship between plethysmographic pressure and Pm. During these respiratory efforts against the closed shutter,
TGV is decompressed and compressed respectively. Because the total amount of gas in
the plethysmograph–lung system is constant, DV causes corresponding changes in
plethysmographic gas pressure during compression and decompression of thoracic gas.
The change in plethysmographic pressure is then measured in terms of the change in
TGV, DV, and denoted shift volume.
With the glottis maintained open, the change in PA during respiratory efforts against a
closed shutter may be measured by recording the change in Pm. In normal subjects, the
change in Pm closely approximates that in PA during panting efforts [1]. However, the
assumption that change in PA can be measured accurately by Pm during panting efforts
against a closed shutter in patients with airflow obstruction has been questioned. Several
groups have reported significant differences between changes in oesophageal pressure
and Pm during panting efforts against a closed shutter in subjects with airflow
obstruction [17–22]. This is discussed more extensively in the sections below:
Pathophysiological manifestations and Measurement of thoracic gas volume.
When only slow (1 Hz) panting efforts against a closed shutter are utilised [17–22], it is
possible to measure changes in PA from Pm. Pm (PA) pressure is plotted against
simultaneous plethysmographic pressure changes (measured as the shift volume) during
respiratory efforts against a closed shutter to measure absolute TGV.
The measurement of TGV is summarised by the following equations, for small changes
in pressure and volume.
Boyle~Mariotte0 s Law : P:V ~constant under isothermal conditions
ð1Þ
During airway occlusion, usually at resting end-expiration, the following equations
describe TGV and PA. The inspiratory or expiratory effort against the closed shutter
will decrease or increase PA by DPA, and increase or decrease TGV by a small volume
change, DV.
PA:TGV~(PA{DPA)(TGVzDV )
ð2Þ
Expanding and rearranging equation (2).
TGV~(DV =DPA):(PA{DPA)
ð3Þ
Since DPA is very small compared to PA (v2%) it is usually omitted in the differential
21
M.D. GOLDMAN ET AL.
term.
TGV*(DV =DPA):PA with PA~Pbar{PH2 O,sat
ð4Þ
TGV*(DV =DPA):(Pbar{PH2 O,sat)
ð5Þ
As noted above, during the respiratory efforts against the closed shutter, the change
in PA i.e. DPA, is recorded as DPm. DV, the shift volume, is measured by the
calibrated plethysmographic gas pressure transducer.
In whole-body plethysmographs, where sRaw is measured during shallow panting,
TGV is determined at a lung volume that is the most comfortable for the patient. This
volume is usually greater than resting FRC, because of comfort factors for normal
subjects and because of flow limitation in patients with obstructive lung disease [23, 24].
Accordingly, this volume increment above resting FRC must be subtracted to provide
FRCpleth [9].
The measured TGV additionally includes any apparatus dead spaces (Vd,app) as well as
any volume inspired above resting end-expiratory lung volume at the moment of
occlusion (Vt,occ). Hence FRCpleth can be derived from TGV by subtraction of these two
volume components.
FRCpleth~TGV{V d,app{V t,occ
ð6Þ
In contrast to gas dilution measurements of FRC (FRCdil), FRCpleth includes all
TGV even if some may not be in communication with the airway opening. Thus, the
value of FRCpleth serves as the methodological anchor for determination of absolute
TGV both at residual volume (RV) and total lung capacity (TLC). The measurement
procedure for determining FRCpleth is more complicated than the recording of sRaw
which is described in the next section, because the subject must respond with normal
breathing efforts while ventilation is interrupted by the closed shutter. Therefore, this
manoeuvre requires more subject cooperation, and FRC may vary from test to test.
In contrast, RV and TLC are more fixed and may be determined immediately after
measurement of FRC by slow exhalation to RV followed by inhalation to TLC.
During tidal breathing, the shutter mechanism is activated by the operator using
computer control, and closes at the end of the following tidal expiration. Most plethysmographs program shutter reopening at a predetermined maximal occlusion time or after
the subject has generated predetermined cumulative inspiratory and expiratory Pm changes
against the shutter or a number of zero-pressure crossings. These different criteria were
introduced to create reliable test results with optimal comfort for the subject. Subjects
should always be informed to remove the mouthpiece from their mouth in the event the
shutter does not open or if the subject senses substantial difficulty breathing.
During the occlusion phase the subject is asked to continue normal breathing efforts
against the closed shutter. To optimise measurement quality, shutter closure settings are
selected to allow recording of at least one positive and one negative Pm change when the
shutter is closed. Plethysmographic shift volume and the corresponding Pm changes are
displayed on an X–Y graph as shown in figure 4.
As in all pulmonary function evaluations, it is recommended that three replicates of
the measurement of TGV are recorded and saved. Quality of the measurement is
reflected, in part, by the variability of replicate trials. Quanjer et al. [25] suggests a
maximal deviation of 5% between the individual trials.
Determination of specific resistance
During assessment of sRaw, it is emphasised that the relationship between airflow and
shift volume, initially described by DuBois et al. [2] does not define Raw. Raw is defined
22
WHOLE-BODY PLETHYSMOGRAPHY
3
2
Pm kPa
1
0
-40
-20
0
20
40
-1
-2
Shift volume mL
-3
Fig. 4. – Respiratory effort against closed shutter in the same patient as in figures 5–7, showing mouth pressure
(Pm) plotted on vertical axis and shift volume (DV) on the horizontal axis. Expiratory effort results in positive
Pm and negative DV, and vice-versa for inspiratory effort. The slope of Pm versus DV is proportional to the
functional residual capacity determined by plethysmography (FRCpleth; 3.8 L). This tracing shows good
coordination of the obstructed inspiratory and expiratory efforts, with only small departures from a single line.
only by combining the measurement of sRaw and the TGV measurement during occluded
respiratory efforts [1, 2].
Mouth flow during spontaneous breathing is continuously recorded from the
pneumotachygraph and displayed on a graphic X–Y display versus the shift volume
produced by thoracic compression and decompression as shown in figure 5. As noted
above, the shift volume, DV, excludes lung volume change due to gas flow in and out of
the lung.
Thermal and humidity effects arise during inspiration of plethysmographic air and
subsequent expiration of warm humid alveolar air. Electronic compensation for thermal
and humidity effects was introduced to permit tidal breathing [26]; however, current
whole-body plethysmographs commonly incorporate algorithms to compensate for
thermal/humidity effects so that the graphic X–Y display of the sRaw loop is closed as
2
Flow L·s-1
1
0
-40
-20
0
20
40
-1
Shift volume mL
-2
Fig. 5. – Tracing of a specific resistance (sRaw) loop in a patient with airflow obstruction, showing the slope
used for calculation of total sRaw (sRtot; 3 kPa?s). Mouth airflow (V9) is plotted on the vertical axis with
inspiratory flows positive and expiratory flows negative. Shift volume (DV) is plotted on the horizontal axis with
inspiratory shift volumes positive and those during expiration negative. See text for discussion.
23
M.D. GOLDMAN ET AL.
completely as possible during inspiration, during tidal breathing, without the need for
rapid shallow respirations.
Equation 7 summarises the relationship between mouth flow, V9, measured by the flow
meter and simultaneously measured plethysmograph pressure, calibrated in terms of shift
volume, DV, to derive sRaw.
sRaw~(DV =V 0 ):(Pbar{PH2 O,sat)
ð7Þ
sRaw is thus determined as the product of dry gas Pbar and the ratio of shift volume to
mouth flow.
It must be emphasised that the commonly utilised slope of the X–Y display of the sRaw
loops in this measuring step does not directly represent Raw (i.e. pressure–flow loops) as
often assumed, but is instead, sRaw. Thus, the slope does not yield a value for Raw, but
requires knowledge of TGV prior to calculation Raw. The sRaw loop is influenced by Raw
and TGV and its inclination rotates clockwise if either Raw or TGV or both are increased.
Acquisition of sRaw data in a whole-body plethysmographic measurement requires
little cooperation from the subject, as this is commonly done in current plethysmographs
during tidal breathing, rather than using the voluntary rapid shallow panting method
originally reported by DuBois et al. [2].
As in all plethysmographic applications, subjects should sit upright and avoid neck
flexion or rotation. After adapting to the measuring conditions during tidal breathing
through the pneumotachygraph, it is recommended that at least 5–10 sRaw loops should
be recorded as one trial. Normally, three replicate trials are recorded and saved. Optimal
quality of the recording is achieved when sRaw loops are regular and reproducible with
the loop nearly entirely closed, although patients with significant airflow obstruction
manifest open loops during expiration.
Numerical parameters calculated from the specific resistance loop. The content of the
sRaw loop is often quite complex and is not a simple narrow linear oval loop, especially in
the presence of peripheral airway disease, as initially described by DuBois et al. [2]. Since
the sRaw loop includes varying flows throughout the tidal breathing respiratory cycle,
different investigators have utilised different portions of the loop to approximate a
"representative" value for the entire cycle.
The total specific resistance (sRtot) [27] and effective specific resistance (sReff) [23] have
been well established and both are utilised in clinical laboratories. These approaches,
along with use of the linear portion of the sRaw loop between inspiratory and expiratory
flow rates of 0.5 L?s-1 [28, 29] are designed to provide a linear approximation of sRaw.
Such linear approximations are generally comparable in patients with normal respiratory
mechanics, but all of these approaches manifest interpretative compromises in advanced
obstructive lung disease. The specific characteristics of these different approaches are
discussed in a subsequent section with physiological interpretation.
Total specific resistance. The sRtot, as described by Islam and Ulmer [27], is determined
by a straight line between maximal inspiratory and maximal expiratory shift volume
points as shown in figure 5.
The outstanding characteristic of sRtot is its sensitivity to partial obstruction of
peripheral airways. The potential disadvantage of sRtot would appear to be a greater
variability from test to test, as a consequence of using only two points at the extremes of
inspiratory and expiratory shift volume.
Effective specific resistance. sReff, as introduced by Matthys and Orth [23], extended
the dimensional analysis applied by Jaeger and Otis [30] to integrate effects of variable
24
WHOLE-BODY PLETHYSMOGRAPHY
flows and nonlinearities of mouth flow-shift volume loops during tidal breathing. They
calculated sReff during tidal breathing from the quotient of the integrated shift volume–
volume loop (flow resistive work of breathing) and the integrated flow–volume loop
(fig. 6a). This ratio defines the slope of a line that represents sReff. Figure 6b shows the
placement of this line within the sRaw loop, defined by performing a least-squared fit of
the line of Matthys and Orth [23] to the points that make up the sRaw loop.
The outstanding characteristic of sReff is its reflection of an integrative assessment of
airway behaviour throughout the entire tidal breath. Digital integration of the respective
loops shown in figure 6a improves the signal-to-noise ratio. sReff reflects larger central
airways somewhat more prominently than sRtot.
Specific resistance at 0.5 L?s-1. DuBois et al. [2] initially measured the slope of the sRaw
loop at a defined fixed flow of 1 L?s-1, noting small increases of the slope in normal
Flow L·s-1
a)
V´I
-DV
+DV
Shift volume mL
V´E
sReff
Volume L
Volume L
VT
-DV
AE
DV0
BE
AI
+DV
V´E
BI
V´I
Flow L·s-1
Shift volume mL
FRCpleth
b)
2
Flow L·s-1
1
0
-40
-20
0
20
40
-1
Shift volume mL
-2
Fig. 6. – a) Schematic representation of graphic integration of flow (V9), volume (V), and shift volume (DV)
parameters, adapted from Matthys and Orth [23], showing flow-resistive work of breathing during inspiration
and expiration (AI and AE, respectively) and flow–volume loop during tidal breathing (BI and BE), used in the
calculation of effective specific resistance (sReff). See text for discussion. b) Same tracing of specific resistance (sRaw)
loop as in figure 5, but showing the slope resulting from calculation of area ratio shown in figure 6a, equal to sReff
positioned on the sRaw loop by regression technique (sReff = 2.7 kPa?s). VT: tidal volume; V9E: expiratory flow;
V9I: inspiratory flow; FRCpleth: plethysmographic functional residual capacity. See text for discussion.
25
M.D. GOLDMAN ET AL.
subjects at 0.75 and 0.5 L?s-1. Subsequently, the flow range has commonly been limited to
the relatively linear portion of the sRaw loop between inspiratory and expiratory flow rates
of 0.5 L?s-1 [28, 29] for definition of sR0.5, as shown in figure 7.
The potential advantage of sR0.5 is that it standardises the flow at which resistance is
measured. In normal subjects, but particularly in patients with airflow obstruction,
resistance is dependent upon flow rate, so this approach offers less inter-individual
variability. The parameter sR0.5 reflects primarily the behaviour of larger, more proximal
airways, with much less sensitivity to peripheral airway abnormalities.
Specific conductance. The reciprocal of sRaw is denoted specific conductance (sGaw).
sGaw~1=sRaw
ð8Þ
When calculating sGaw, it must be defined with respect to which calculation of sRaw is
performed, according to the definitions listed in the sections above. The conversion of
sRaw to sGaw is not simply a mathematical procedure, but is based on the original
observations of Briscoe and DuBois [31] that the major determinant of Raw in normal
subjects is lung volume and, accordingly, that the relationship between lung volume
and conductance is linear within and between individuals. Thus, sGaw is a "volumenormalised" expression for airway conductance.
Calculation of airway resistance and conductance
Finally, the commonly used clinical parameters of body plethysmography, Raw and
Gaw, are calculated using sRaw and corresponding TGV, as defined below:
ðÞ
Raw~sRaw=TGV
Or corrected for the average lung volume during tidal breathing, where VT represents
tidal breathing.
ð9Þ
Raw~sRaw=(FRCplethzVT =2) G aw~1=Raw
In practice, the measurements of TGV are conveniently performed immediately after
the sRaw breathing loops; and three replicates are recommended. Quality of the
measurement is reflected in part by the variability of replicate trials and, in part, by
2
Flow L·s-1
1
0
+0.5 L·s-1
-40
-20
0
20
40
-0.5 L·s-1
-1
Shift volume mL
-2
Fig. 7. – Same tracing of specific resistance loop as in figure 5, showing the slope used for calculation of specific
resistance 0.5 (sR0.5; 2.5 kPa?s). See text for discussion.
26
WHOLE-BODY PLETHYSMOGRAPHY
how closely the Pm – plethysmograph pressure tracing approximates a straight line.
A good quality tracing is shown in figure 4, where departures from the computer
regression line are very small over a wide range of Pm.
By definition, inaccuracy in the determination of TGV or FRCpleth will cause a
proportional error in the estimation of Raw and Gaw. For this reason, and because it is
technically more demanding for patients with airflow obstruction to make respiratory
efforts against a closed shutter than for tidal breathing, some clinicians restrict their most
careful attention to sRaw and sGaw [23, 27, 30, 32]. Additionally, in many patients with
COPD, Raw appears to be nearly within normal limits, due to manifest compensatory
lung hyperinflation, especially when measured between 0.5 L?s-1 inspiratory and
expiratory flow. In these cases, sRaw and sGaw still show abnormality, because of the
increased TGV maintained during tidal breathing. In a subsequent section, an alternative
approach to estimation of TGV during tidal breathing only is described, avoiding
voluntary respiratory efforts against a closed shutter [24].
Spirometric measurement
It is often convenient to complete a body plethysmographic measurement with
spirometric measurements. Commonly this is done immediately after TGV has been
determined, using a slow exhalation below resting FRC to minimal lung volume, i.e.
performance of an expiratory reserve volume (ERV) effort. This is followed by an
inspiratory vital capacity effort (IVC) to TLC, followed by a maximal forced expiration
for determination of forced expiratory volume in one second (FEV1) and FVC.
In this way, all the primary pulmonary subdivisions can be recorded as absolute gas
volumes. These include TLC, FRCpleth and RV. RV may be calculated by subtracting
ERV from FRCpleth.
RV~FRCpleth{ERV
ð10Þ
TLC is determined by adding the maximal VC recorded, usually IVC, to RV.
TLC~RVzIVC
ð11Þ
Inspiratory capacity (IC) is the difference between TLC and FRCpleth.
IC~TLC{FRCpleth
ð12Þ
The spirometric data described above are conveniently recorded from the flow meter
in the whole-body plethysmograph. Issues relevant to spirometry are reviewed and
discussed in another chapter of this Monograph. However, it is relevant to note here
that, using the "transmural" pressure-compensated integrated-flow plethysmograph
it is possible to view the maximal expiratory flow–volume (MEFV) curve with respect
to volume displacements of the thorax, including those due to compression, during
forced expiration [12]. This is a more reliable method of assessing for the presence of
expiratory-flow limitation during resting breathing compared with maximal forced
expiration, than spirometry using only integrated mouth flow as the volume axis.
Using such a transmural plethysmograph, it is immediately evident that the VC
measured from thoracic wall displacements is greater than that measured from
integrated flow, because of compression of thoracic gas trapped behind closed small
airways at low lung volumes. While this is not important in making clinical decisions,
the clinical value of thoracic wall displacements during tidal breathing is a significant
issue in patients with chronic airflow obstruction and is discussed below (section
Clinical utility of whole-body plethysmography).
27
M.D. GOLDMAN ET AL.
Pulmonary function using whole-body plethysmography
Initial interpretation of body plethysmographic parameters usually considers
measured values in comparison to established normative data. However, it is often
preferable to use the patient as his own control, by assessing the trend of measurements
over time or to repeat measurements after therapeutic challenge. Additionally,
plethysmography may be repeated after bronchial challenge to assess airway reactivity.
Predicted and limit values for airway resistance
Relatively few studies have established predicted values of Raw in adults. Age
differences have relatively unimportant effects, as first noted by Briscoe and DuBois
[31]. Ulmer and coworkers [33, 34] reported an average Rtot for healthy adults of
0.22 kPa?s?L-1 and defined an upper limit of normal Rtot as 0.35 kPa?s?L-1. Matthys et
al. [35] introduced normative equations for sRtot and sReff, and reported an average¡sd
value for Reff of 0.2¡0.0967 kPa?s?L-1.
Recently Van der Velden et al. [36] compared Rtot, Reff and R0.5 in 78 healthy adults
with average¡sd values for Rtot of 0.19¡0.07 kPa?s?L-1, for Reff of 0.15¡0.06 kPa?s?L-1
and R0.5 of 0.13¡0.05 kPa?s?L-1. These comparative values are useful for current
guidance. Quanjer [37] tabulated data in 1983, including a large 1970 study of Raw
during tidal breathing, in both males and females. He selected an upper limit of normal of
0.3 kPa?s?L-1 for both males and females. However, the age of these data argue for the
value of undertaking new studies of normative values for Raw, sGaw and, possibly,
absolute lung volumes using modern plethysmographs with thermal/humidity effects
compensated by numerical algorithms.
In younger children, Klug and Bisgaard [38] have measured sRaw with the child
accompanied by an adult within the plethysmograph. As expected with growth and
increase in lung size, Raw decreases with age in children. Predicted values for children
have been reported by Zapletal et al. [39].
Predicted value for thoracic gas volumes
Body size and lung size in adults may vary according to ethnic origin and some
normative values have been reported by different authors to correspond to populations
served in their communities. Quanjer et al. [25] reported standardised values for FRC,
RV and TLC with spirometry in adults and Ulmer et al. [34] reported standardised
values for TGV and FRCpleth in adults. Zapletal et al. [39] reported plethysmographic
volumes for children.
Assessment of bronchial reactivity
Measurement of sRaw (sGaw) has been used clinically for assessment of bronchial
responsivity. Because sRaw and sGaw are commonly measured during tidal breathing, they are influenced both by Raw as well as changes in resting lung volumes (FRC).
Since both resistance and resting end-expiratory lung volume may change during
bronchial or therapeutic challenge, sRaw and sGaw provide useful practical assessments of
airway responsivity, even in the absence of a determination of absolute TGV. In adults
or children unable to perform the measurement of TGV, sRaw or sGaw provides useful
clinical guidance, although American Thoracic Society and European Respiratory
Society guidelines suggest separate documentation of Raw and changes in FRC. Such
28
WHOLE-BODY PLETHYSMOGRAPHY
measures of airway response during tidal breathing are often considered preferable to
spirometric assessments [40]. The commonly used limit for bronchoprovocation is a 15 or
20% decrease in FEV1 relative to control baseline FEV1. The comparable limit for
sRtot is 100%, for Rtot 50% increase and for sGtot 40% decrease from baseline,
respectively [40].
Therapeutic challenge may be similarly compared to baseline sRaw and judged by the
degree of reversibility, whether limited in magnitude (partial reversibility) or more
complete, such that sRaw values reach the normal range. Reversibility, whether partial or
complete, can be assessed as the improvement quantified as per cent of the predicted
value.
Interpretations of whole-body plethysmography
Pathophysiological manifestations
While the numerical values of sRaw and sGaw, Raw, Gaw, and TGV may be compared
with normative data where they are available, and for assessment of bronchial and
therapeutic challenge, the linear approximations used to derive numerical values provide
a limited capacity for the understanding of pathophysiology. Further physiological
interpretative information is available from the shape of the sRaw loops. The additional
value of these graphic displays is analogous to the additional value of the flow–volume
curve, relative to simple numerical values of FEV1 and FVC.
The infrastructure of physiological interpretation of sRaw loops is the relationship
between airflow measured at the mouth and shift volume (V9 versus DV). Shift volume
represents the volume changes in TGV that occur during compression and decompression of thoracic gas, not including the volume changes due to airflow in and out of the
lung, and this shift volume is related to airflow resistance. When airflow resistance is the
dominant contribution to shift volume, changes in PA and shift volume usually manifest
a linear relationship to airflow at the mouth. This is made use of in estimates of sRaw
between the limits of 0.5 L?s-1 inspiratory and expiratory flows. However, even in normal
subjects when airflow rate is substantially larger than 0.5 L?s-1 it is common to observe
slight alinearity of sRaw, as noted in original report of DuBois et al. [2].
Mild obstructive lung disease may manifest as only minimal nonlinearity of sRaw
loops. However, in advanced obstructive lung disease, it is now well known that dynamic
compression of intrathoracic airways is associated with disproportionate increases in
intrathoracic pressure relative to airflow. Stanescu et al. [18] and Rodenstein et al. [19]
used oesophageal pressure to estimate pleural pressure during respiratory efforts against
a closed shutter in patients with airflow obstruction These studies demonstrated that, in
the presence of increased airflow resistance, mouth occlusion pressure changes
underestimate those of oesophageal (and alveolar) pressure during panting at frequencies
w1 Hz. Other investigators confirmed the inaccuracy of TGV measured during panting
against a closed shutter at frequencies w1 Hz, and suggested that their results were
consistent with nonhomogeneous mechanical properties of airways and lung tissue time
constants [17, 22]. Furthermore, in patients with severe airflow obstruction, there may be
areas of the lungs that do not communicate with central airways, and, therefore, do not
ventilate during tidal breathing, as evidenced by measures of "closing volume" that occur
at lung volumes that may exceed FRC [41, 42].
Islam and Ulmer [27] provided a comprehensive evaluation of effects of airway
closure using plethysmographic measures of the altered relationship between changes in
intrathoracic pressure relative to airflow. They reasoned that the marked narrowing or
29
M.D. GOLDMAN ET AL.
closure of small airways that occurred at low lung volumes, defined as closing volume
[41, 42], should cause an abrupt decrease in plethysmographic gas pressure. They plotted
apparent Rtot as a function of lung volume, and showed a dramatic increase in apparent
Rtot in patients with airflow obstruction at low lung volumes, manifest to a lesser degree
in normal subjects [27]. In normal subjects, they determined closing volume at lung
volumes below FRC (i.e., within the ERV), which they associated with a significant
increase in apparent Rtot. In patients with chronic airflow obstruction, they were unable
to determine a closing volume because of technical limitations; however, they measured a
substantial increase in apparent Rtot within the IC. These authors utilised changes in
apparent Rtot as a reflection of compression of gas in nonventilated airspaces. The
clinical implication of such changes is discussed below (section Extending the clinical
utility of whole-body plethysmography).
Shortly after the report of Islam and Ulmer [27], Matthys and Orth [23] described
the contribution of these pathophysiological disturbances to a dissociation between
maximal shift volume and maximal flow. They extended the dimensional analysis applied
by Jaeger and Otis [30] to integrate these contributions to an "effective resistance" that
included the effects of the entire range of variable flows during tidal breathing and
nonlinearities in the sRaw loop. They measured the areas of graphic plots of shift volume
versus volume and of flow versus volume during tidal breathing, determined
planimetrically during playback of plethysmographic signals recorded on magnetic
tape (fig. 6a). They divided the integrated shift volume–volume loop (the flow resistive
work of breathing, A in fig 6a) by the flow–volume loop (B in fig 6a) to derive sReff. They
calculated effective resistance from the quotient of sReff and mean ventilated lung volume
(FRCpleth z VT/2).
Reff ~½(A=B):(Pbar{PH2 O,sat)=(FRCplethzV T=2)
ð13Þ
Matthys and Orth [23] performed these calculations from analysis of signals
recorded on magnetic tape; but this is now readily calculated by digital algorithms in
modern computer-assisted plethysmographs. Despite the obvious attraction of an
integrative approach, such as that of Matthys and Orth, the analysis and
interpretation of multiple graphic displays, including flow–volume loops, shift
volume versus volume and shift volume versus flow loops, is not feasible in the clinical
pulmonary function laboratory. Accordingly, calculation of the numerical value of
sReff is done by computer algorithm, and the resulting slope is positioned within the
conventional sRaw loop using regression techniques. In this way, sReff can be
compared conveniently to sRtot and sR0.5 if desired [36].
Since the contributions of dynamic compression of intrathoracic airways and
compression of nonventilated lung areas make the sRaw loops highly nonlinear and
contribute to characteristic shapes of the shift volume versus mouth flow X–Y graph
displayed in current body plethysmographs, these characteristic shapes are now discussed
in detail.
Characteristic specific resistance loops
Characteristic sRaw loops are shown in figure 8. The tracing labelled a) in figure 8
displays a schematic sRaw loop in a normal subject during tidal breathing, which is shown
after numerical software compensations to close the sRaw loop. Normal subjects manifest
a steep linear loop during tidal breathing without hysteresis. In contrast, during
voluntary panting efforts, the upper and lower end portions of the loop may become
slightly curvilinear. The curvilinearity is in the form of a very slight "S" shape, analogous
to that shown in tracing d), but much less exaggerated. In normal subjects during
30
WHOLE-BODY PLETHYSMOGRAPHY
2
a)
b)
d)
1
Flow L·s-1
c)
0
-1
Shift volume mL
-2
Fig. 8. – Schematic representation of specific resistance loops in a) a normal subject, b) a subject with increased
large airway resistance, c) a subject with chronic airflow obstruction d) and a subject with upper airway
obstruction. Mouth flow (V9) is plotted on the vertical axis, with inspiration positive and expiration negative.
Shift volume is plotted on the horizontal axis. See text for discussion.
voluntary panting, the flattening of the sRaw loop at the upper right extremity (midinspiration) and at the lower left extremity (mid-expiration) of the loop are only barely
visible, depending on the absolute value of flow rates achieved. While the accepted
numerical limits of normality are broad, it is the characteristic shape of the sRaw loop,
immediately apparent from direct observation, that guides clinical interpretation.
Tracing b) in figure 8 is typical of subjects with large (central) airway constriction
that is relatively uniform (and not a localised stenosis) and without significant small
airway obstruction. This might be seen in a patient with mild asthma. Here a linear sRaw
loop that is tilted clockwise, manifesting a slope less steep than normal, reflects increased
Raw.
In subjects with normal pulmonary mechanics or uniformly increased large airway
constriction, as noted immediately above, the sRaw loop has little or no hysteresis
("openness" of the loop). In patients with nonhomogeneous small airway partial
obstruction, the sRaw loop manifests the characteristic shape shown by tracing c) in
figure 8. The loop is quite open, especially during expiratory flow. A large shift volume
appears at mid expiration, without corresponding increases in expiratory flow. Such
alinearities may represent expiratory flow limitation and/or dynamic airway compression. It is well known that expiratory flow limitation and dynamic airway compression
may occur during tidal breathing in COPD [41, 42], and this contributes to the
characteristic shape of the sRaw loop in tracing c).
Compression of nonventilated airspaces will also contribute to the leftward
displacement of the shift volume versus mouth flow tracing. Rodenstein et al. [43]
obstructed right lung middle and lower lobes in normal humans to assess the effect of
nonventilated airspace on measurement of TGV. They did not report sRaw loops during
such obstruction, but the similar TGV reported without and with nonventilated airspaces
implies, by definition, that compression of TGV with increases in PA is quantitatively the
same without and with nonventilated airspaces. Thus, the relationship between shift
volume and airflow will be distorted as in tracing c) by compression of nonventilated
airspaces. Changes in plethysmographic volume with compression of gas behind closed
airways were demonstrated by Davis et al. [13] and, as noted above, by Islam and Ulmer
[27] prior to that.
It may be seen from the shape and direction of tracing c), comparing early expiratory
31
M.D. GOLDMAN ET AL.
flow with late expiratory flow at the same value of mouth flow, that shift volume is less
early in the expiration compared to late in expiration at the same flow. This hysteresis
defines a nonlinear relationship of shift volume to mouth flow that may include
contributions of dynamic airway compression and compression of nonventilating
airspaces to the overall TGV compression during expiration. The single lines drawn in
figures 5–7 represent lines defined as sRtot, sReff and sR0.5. It is readily apparent that such
a single line drawn for sRtot reflects a single index that includes important nonlinearities
occurring during expiratory airflow. This single line is very different from a
"representative" line that might be drawn during inspiratory airflow only or the line
corresponding to sR0.5 in tracing c). More important than any attempt to quantify the
complex shape of the sRaw loop by a single index, the X–Y display itself reveals the highly
abnormal mechanical behaviour during expiratory airflow in tracing c). These
abnormalities include contributions from nonlinear expiratory airflow resistance,
dynamic airway compression and compression of nonventilated airspace. The latter
two factors contribute to the increased shift volume late in expiration compared to early
in expiration, even at an identical mouth flow.
Numerical analysis of tracings in patients similar to those in figure 8c, after dividing by
TGV, may be compared with normative values listed above in the section Predicted and
limit values for airway resistance. It should be noted that calculation of measured values
as per cent predicted may differ in plethysmographs available from different
manufacturers. Such calculations should state whether "predicted" is the mean expected
value or the upper limit (for resistance) of accepted normal values. Equally importantly,
extension of the study of Van der Velden et al. [36] should be undertaken with modern
commercially available plethysmographs to confirm their predicted values, including a
larger normal population sample and to compare Rtot, Reff and R0.5 in patients with
chronic airflow obstruction done at baseline and following therapeutic challenge. The
value of such extensions of plethysmography is discussed below in the section Extending
the clinical utility of whole-body plethysmography.
Because of mechanical nonhomogenities in the lung and airways in obstructive lung
disease, it is not entirely satisfactory to attempt to summarise Raw by a single number.
Future clinical investigations might usefully include discrimination between inspiratory
and expiratory Reff, to recognise the predominance of abnormality during expiratory
airflow. An alternative distinction can be made by looking at the parameter most
commonly used in North America, Raw between inspiratory and expiratory flow rates of
0.5 L?s-1. It can be seen in tracing c) that the line corresponding to sR0.5 would be
substantially steeper (less abnormal) than that for sRtot. This reflects, in part, the smaller
flow rates, higher lung volume and lack of dynamic airway compression during late
inspiration/early expiration. It is fair to state that inspection of the shape of the sRaw loop
displayed as the X–Y graph is equally useful diagnostically as any single or combination
of numerical values.
Tracing d) in figure 8 shows the influence of a fixed or functional stenosis of
the upper airways, for example laryngeal abnormality, or paralysis of one vocal cord.
This type of "orifice" constriction manifests flow limitation during inspiration, such
that, at sufficiently high flows, further increases in driving pressure do not result in any
increase in airflow. This reflects localised upper airway obstruction, analogous to that
which pertains in the maximal expiratory flow–volume curve. Thus, during forced
expiration, when a critical driving pressure for expiratory airflow (intrapleural pressure
for forced expiration) is achieved, further increases in driving pressure do not cause
any further increases in flow rate. A similar flow limitation may occur in the
extrathoracic airway during inspiration, as shown in the upper right portion of tracing d)
in figure 8.
32
WHOLE-BODY PLETHYSMOGRAPHY
Clinical utility of whole-body plethysmography
The utility of whole-body plethysmography is discussed from the perspective of clinical
respiratory medicine by Brusasco and Pellegrino [44] and physiological considerations
are presented in detail by Pride and Macklem [45].
Measurement of thoracic gas volume
The raison d’etre of whole-body plethysmography is the measurement of lung volumes.
Accordingly, the first acknowledged clinical benefit of body plethysmography is the
definition of restrictive lung disease [46]. Normative data for TGV and pulmonary
subdivisions allow definition of restrictive lung disease as distinct from obstructive, in the
presence of a reduced VC. Definition of abnormally increased lung volumes in
obstructive lung disease is a further appropriate clinical use of whole-body
plethysmography. While lung volumes can be measured by gas dilution techniques, it
is well known that dilution techniques measure only the volume of ventilated airspaces.
Accordingly, when whole-body plethysmography is combined with dilution measures of
lung volumes, the volume of trapped gas is estimated by the difference between FRCpleth
and dilutional FRCHe. Because FRC varies to some degree from breath to breath, a
further comparison of calculated RV determined with dilution and plethysmography
provides useful information concerning trapped gas.
The voluntary rapid shallow obstructed respiratory efforts described by DuBois et al.
[1] appear to permit equilibration of intrathoracic gas and Pm, and, accordingly, a
realistic estimate of changes in PA from Pm measurements in normal subjects. However,
in the presence of intrathoracic airway obstruction, rapid obstructed panting efforts
overestimate TGV because the change in Pm underestimates the change in PA [18, 19].
Stanescu et al. [18] and Rodenstein et al. [19] investigated normal and asthmatic
subjects. They compared changes in mouth pressure with those of oesophageal pressure
during obstructed panting efforts and showed that in normal subjects Pm and
oesophageal pressures during panting efforts against a closed shutter were comparable.
However, in the presence of airflow obstruction, changes in Pm significantly
underestimated those in the oesophagus, taken to be equal to PA changes during
respiratory efforts against a closed shutter. Airway obstructions were either diffuse, as in
asthmatic subjects, or in the lower trachea, induced in normal subjects by inflating a
balloon in the lower trachea. This group then bypassed the upper airways with a cuffed
endotracheal tube and showed comparable occlusion pressure changes between the
endotracheal tube opening and the oesophagus. They concluded that an increased degree
of airflow obstruction, increased compliance of the upper extrathoracic airways and
increased rate of panting all combine to cause the underestimation of PA change by Pm,
and consequent overestimation of TGV. This work and that of others [17, 20–22] resulted
in a recommendation of panting at 1 Hz to optimise the measurement of TGV.
Thus, the assumption implicit in the original work of DuBois et al. [1] by use of
changes in Pm to represent changes in PA during panting efforts against a closed shutter
has been demonstrated to be unwarranted in patients with significant airflow
obstruction, unless very slow panting efforts are performed. However, such slow
panting efforts require considerable coordination on the part of the patient, and, in
practice, tidal breathing is much more reliably assessed in current commercial
plethysmographs with the aid of computer-assisted compensation for thermal and
humidity effects.
A second assumption is that the changes in body volume during panting efforts against
a closed shutter are essentially only those of TGV. This assumption has been reinforced
33
M.D. GOLDMAN ET AL.
by Brown et al. [47], who investigated the effects of panting efforts at different volumes
within the VC and with different amounts of abdominal air introduced into the stomach
via a nasogastric catheter. They reported that when panting efforts were performed near
RV and near TLC, discrepancies of 3–5% of true TLC could be related to abdominal gas
volume, but when panting efforts were performed at FRC the effect of abdominal gas
volume on measurement of TGV was negligible. The simplest form of Boyle-Mariotte’s
Law used in manual calculations of TGV [1] has been evaluated by Coates et al. [11] who
included calculation of TGV using the complete Boyle-Mariotte’s law equation
(Equation 3, section Determination of thoracic gas volume and functional residual
capacity) and demonstrated errors in the order of ¡3% during panting and ¡2–9% during
a single inspiratory effort against a closed shutter as recommended for children [48].
Although such discrepancies are not likely to influence clinical decisions, the authors
argued that they are easily avoided using modern computational methods in automated
whole-body plethysmographs [49].
The foregoing analysis and review of efforts to optimise the measurement of FRCpleth
has emphasised the cooperation required of the patient, including panting efforts against
a closed shutter at a controlled low frequency in addition to maintenance of an open
glottis during obstructed respiratory efforts. These constraints prompted Agrawal and
Agrawal [24] to measure TGV during tidal breathing without obstructed respiratory
efforts. These authors reasoned that since sRaw is expressed numerically by the product
of TGV and Raw, addition of a known resistance in the respiratory path would permit
determination of TGV by subtraction. Thus:
sRaw1~Raw:TGV and
ð14Þ
sRaw2~(RawzRadd):TGV
Subtracting Equation 14 from Equation 15 yields:
sRaw2{sRaw1~Radd:TGV; and TGV~(sRaw2{sRaw1)=Radd
ð15Þ
ð16Þ
It is implicit in these equations that TGV must be constant between tidal breathing without and with the added resistance, airflow at which sRaw is measured
must be the same without and with the added resistance; and that airway mechanics
can be modelled as a linear system. These authors measured sRaw manually from
an oscilloscope screen at the onset of inspiration up to 0.5 L?s-1 inspiratory flow
without and with added resistance. The added resistance was brought into the
respiratory path by a shutter valve which permitted replicate measures of sRaw with
and without added resistance in a constant-volume plethysmograph. Any change in
FRC associated with switching of the shutter valve could be measured by integrated
airflow; however, no changes in FRC were observed. Thus, the first two assumptions
are warranted. The assumption of linear behaviour without and with an added
resistance in front of the mouth may be questioned in patients with airflow
obstruction and nonhomogeneities in lung mechanical properties. Accordingly, the
authors measured TGV during tidal breathing without and with the added resistance,
and compared these results with TGV measured during panting at FRC. Good
agreement between the two methods was obtained in normal subjects and a limited
number of asthmatic and COPD patients in whom baseline Raw ranged from 0.1–
1.5 kPa?s?L-1 [24].
The advantage of estimating lung volume in this manner is that tidal breathing only is
required. Agrawal and Agrawal [24] measured sRaw manually, and it remains to be
determined whether modern computer-assisted plethysmographs will provide comparable FRCpleth results during respiratory efforts against a closed shutter and during tidal
breathing without and with added resistance. It is the authors’ opinion that this approach
34
WHOLE-BODY PLETHYSMOGRAPHY
is worthy of further investigation as it presents a convenient approach to the measurement of TGV that is likely to be more easily applicable to a wide variety of patients.
Measurement of airway resistance
Measures of Raw made in a whole-body plethysmograph demand the constraints and
linear approximations described in previous sections. Accordingly, a single number
defining "resistance" is not entirely satisfactory in patients with substantial airflow
obstruction. Nonhomogeneous lung mechanical properties, expiratory flow limitation and
airway closure all contribute to the highly nonlinear shapes of the sRaw loops described in
previous sections. Limitations of interpretation imposed by the linear approximations
described in sections Numerical parameters calculated from the specific resistance loop,
Pathophysiological manifestations, and Characteristic specific resistance loops point to the
clinical utility of direct visual inspection of the sRaw loops themselves. In addition to
calculating resistance by any of the alternative linear approximations, the shape of the sRaw
loop provides improved understanding of the patients’ pathophysiology.
Plethysmographic sRaw can be measured both during rapid shallow breathing
(panting) and during tidal breathing. The initial description of sRaw [2] utilised rapid
shallow breathing to minimise thermal effects. This had the added advantage of resulting
in full opening of the vocal cords [50]. However, one disadvantage of panting respirations
is that they are almost invariably performed at lung volumes significantly larger than
resting FRC [23, 24], necessitating further corrections to optimise accuracy [9, 11].
Furthermore, controlling panting frequency at a rate of 1 Hz [17–22], as well as requiring
substantial coordination of the patient’s respiratory efforts, also increases the likelihood
of variable glottic opening [50]. Krell et al. [32] demonstrated that quiet breathing sRaw
was equivalent to that obtained during panting. Subsequently, with improved computerassisted compensation algorithms [49], it was possible to program commercial wholebody plethysmographs to measure sRaw during tidal breathing, at normal resting FRC.
Pulmonary resistance, including Raw and tissue viscance, is also available during tidal
breathing from the measurement of oesophageal pressure, although this invasive
procedure is both more time consuming and more uncomfortable for the patient.
Respiratory resistance is available during tidal breathing using the method of forced
oscillation, and is described in another chapter in this monograph. Neither pulmonary
resistance nor forced oscillatory resistance has yet achieved the clinical acceptance of
whole-body plethysmography. Interestingly, the forced oscillation technique was first
introduced by DuBois et al. [51] in the same year that this group first published the
plethysmographic measurement of Raw.
The clinical utility of plethysmographic measurements of Raw and sRaw is attested to
by the fact that they have been considered the "gold standard" for decades for assessing
airway function. In patients with significant airflow obstruction, sGaw is commonly
assessed. This permits lung hyperinflation to be taken into account. Normative values are
available for Raw, sRaw, and their reciprocals, Gaw and sGaw [33–36].
The choice of which measure of resistance is clinically most useful varies among
different investigators and in different countries. Some investigators emphasise the
advantage of Rtot because it includes the effects of multiple mechanical abnormalities
associated with advanced peripheral airway obstruction. Against this is the disadvantage
of test-to-test variability, due to its derivation from only two points (maximal inspiratory
and expiratory shift volumes) of the sRaw loop. Other investigators prefer Reff, because it
integrates the entire ranges of flow, shift volume and lung volume of the complete tidal
breath, and may thus be expected to offer less within-individual variability. Others argue
against the perceived advantages of both these approaches to approximate "resistance"
35
M.D. GOLDMAN ET AL.
because of their sensitivity to nonflow-resistive mechanical effects due to compression of
nonventilating air spaces and also sensitivity to dynamic expiratory intrathoracic airway
compression and expiratory flow limitation during tidal breathing. These mechanical
abnormalities, albeit related to pressure dissipation during airflow in patients with
chronic airflow obstruction, are largely excluded from the calculation of R0.5. For these
reasons, most North American clinicians utilise R0.5, which is derived from a
standardised flow range between late inspiration, z0.5 L?s-1, and early expiration,
-0.5 L?s-1, on the sRaw loop (fig. 7). This calculation results in a lower resistance than
either Reff or Rtot because it is minimally affected by dynamic airway compression or
compression of nonventilating airspace. As such, it reflects primarily the resistance in
larger central airways, is relatively insensitive to changes in peripheral airways and
manifests less test-to-test variability within an individual.
The effects of dynamic airway compression and compression of nonventilating
airspaces lead to a dependence of Rtot and Reff on breathing pattern itself, namely the
degree to which patients with chronic airflow obstruction "force" their expiratory effort.
During resting tidal breathing in normal individuals, expiratory airflow is largely, if not
entirely, produced by stored elastic energy in the chest wall. However, even in normal
subjects, Loring and Mead [52] have shown that resting breathing is associated with a
variable degree of active abdominal muscle recruitment. This active expiratory muscle
recruitment is much more marked in patients with chronic airflow obstruction. Such
patients commonly utilise active expiratory muscle effort to aid expiratory airflow and
manifest expiratory flow limitation even during resting tidal breathing [53]. Depending
upon the patient’s unique individual sensation of their breathing, they may contract their
expiratory muscles to a variable degree during resting tidal expiration, and this active
expiration may change variably with therapeutic challenge. The degree of expiratory
muscle effort will directly influence calculated Reff and Rtot because greater efforts cause
greater shift volumes without corresponding increases in expiratory airflow in the
presence of expiratory flow limitation.
It is clear that there are marked differences between "instantaneous" airflow resistance
during inspiration and expiration in patients with chronic airflow obstruction. These
differences may be appreciated graphically by direct visual inspection of the sRaw loop.
They may be appreciated numerically by deriving separate inspiratory and expiratory
values of Reff, again using the integrated areas of shift volume–volume and flow–volume
loops, as denoted AI/AE and BI/BE for inspiration and expiration separately in figure 6a.
Comparable numerical representation of the mechanical abnormalities that occur during
expiration using R0.5 or Rtot is not possible due to the definition of these quantities based
on the sRaw loops. Instead, graphic display of the sRaw loop is required to appreciate the
prominence of such abnormalities during the expiratory phase [2, 54–56].
However, current computer-assisted plethysmography makes it possible to calculate
"instantaneous" values of airflow resistance, provided TGV is known. During breathing
within the constant-volume plethysmograph, airflow resistance in the lung requires small
amounts of compression of thoracic gas during expiration and expansion of thoracic gas
during inspiration, resulting in the "shift volumes" measured by the pressure change in
the plethysmograph. Calculation of Raw requires measures of PA and airflow. During
free breathing, shift volume can be used to record an index of changes in PA, because
shift volume is the product of TGV, and the change in alveolar pressure, DPA, divided by
initial PA. In other words, the fractional change in PA, [DPA]/(Pbar–PH2O,sat), integrated
over TGV causes a change in TGV equal to shift volume, which, in turn, results in
plethysmographic pressure change. In this way, shift volume provides an index of DPA
provided TGV is known. It must be emphasised, however, that plethysmographic
pressure change during breathing is not equal to DPA. It is much smaller in magnitude
than DPA, and reflects the fractional DPA amplified by TGV.
36
WHOLE-BODY PLETHYSMOGRAPHY
The instantaneous relationship between DV, TGV and PA may be written as:
DPA=(Pbar{PH2 O,sat)~DV=TGV
ð17Þ
This is a restatement of Boyle-Mariotte’s law that, under isothermal conditions, the
fractional change in PA is equal to the fractional change in TGV. Equation 17 may be
rearranged as follows:
DPA~(Pbar{PH2 O,sat)(DV =TGV)
ð18Þ
Thus, instantaneous PA during free breathing can be defined using the product of dry
gas Pbar and the ratio of shift volume to TGV, if TGV is known. This is done by the
computer, continuously in time, from measured signals of shift volume, volume and
airflow after respiratory efforts against a closed shutter have been utilised to calculate
TGV. Instantaneous Raw (iRaw) is then defined by the ratio of instantaneous PA to
instantaneous airflow. This computer calculation has only recently been implemented, and provides a convenient display of Raw throughout the tidal breath, except at
end-expiration and end-inspiration, where iRaw is undefined because airflow is zero,
as shown in figure 9.
It should be noted that Raw calculated in this manner includes nonlinearities in flow
resistance and effects of expiratory flow limitation, and also what may be considered by
some to be "inappropriate" attribution of compression of trapped gas to flow resistance.
Expiratory flow limitation contributes variably to apparent Raw as a function of
respiratory effort: the greater the expiratory muscle effort, the larger the calculated
expiratory Raw at a fixed flow rate. Nevertheless, these contributions may be
2.0
iRaw kPa·s·L-1
1.5
1.0
0.5
FRCpleth
0
3
4
TGV L
5
Fig. 9. – Calculated instantaneous airway resistance (iRaw), plotted on the vertical axis against absolute thoracic
gas volume (TGV) on the horizontal axis, in the same patient as represented in figures 4–7. Dashed lines are
extrapolations during times of zero airflow. FRCpleth is shown by the vertical solid line drawn at 3.8 L TGV.
See text for discussion.
37
M.D. GOLDMAN ET AL.
appropriately considered "resistive". Compression of trapped gas during expiration and
decompression during inspiration are not related to airflow per se, but, nevertheless,
contribute to the total dynamic PA burden during breathing. More importantly, the
degree of trapped gas in patients with airflow obstruction is likely to be related much
more prominently to small airway obstruction than to larger more central airways. Thus,
this Raw will be more sensitive to small airway obstruction than R0.5 and Reff.
It may be seen in figure 9 that there is a progressive increase in calculated Raw
throughout expiration, consistent with the known effects of mechanical abnormalities
during expiratory air flow in patients with airflow obstruction. The envelope of values of
Raw in figure 9 includes wide variability of iRaw throughout the course of the tidal breath
in patients with severe airflow obstruction. For comparison, it may be noted that R0.5 will
be approximately equal to the iRaw values just before end-inspiration, while Rtot and Reff
values will fall near the middle of the expiratory iRaw envelope. This representation may
serve as a useful extension of plethysmographic technique, as noted in the section below.
Extending the clinical utility of whole-body plethysmography
This review draws to its conclusion by extending the exploration of clinical
implications of the complexity of the relationship between shift volume and airflow.
As noted above, this complexity has resulted in three different numerical approximations
to measure resistance derived from different linear approximations of the shift volume–
airflow relationship. The limitations of rapid shallow panting efforts have been described
and the resultant improvements offered by tidal breathing in the determination of
resistance in patients with airflow obstruction.
The potential for tidal breathing estimation of TGV by addition of a known resistance
in front of the mouth has been introduced [24], and will await further investigation using
modern computer-assisted plethysmographs that provide numerical compensation for
thermal and humidity effects during tidal breathing. Investigations in patients with
airflow obstruction should include baseline measures and the response to acute
bronchodilation to fully utilise the scope of experimental conditions in which this
approach might be applicable.
Further investigations that extend the work of Van der Velden et al. [36] in patients
with chronic airflow obstruction will provide useful comparisons of the different
numerical approximations to measurement of resistance. Such investigations will usefully
include response to acute bronchodilation with b-agonists on the one hand and shortacting anticholinergics on the other, in the same patients. In this way, the relative
sensitivity to primarily proximal or distal airway bronchodilation of Rtot, Reff and R0.5
can be assessed in patients with airflow obstruction, with bronchodilator effects in
primarily proximal or distal airways.
The relationship between calculated iRaw and lung volume during the tidal breath has
recently been demonstrated (section above, Measurement of airway resistance). Extension of these studies may prove to be a useful representation for clinicians, permitting a
graphic impression of change in apparent iRaw within the tidal breath. Further
investigations are necessary to define the relative sensitivity of iRaw and expiratory Reff to
interventions that affect primarily larger proximal or smaller peripheral airways.
The relationship between shift volume and lung volume itself is now considered. As
noted in section Spirometric measurement, the VC measured plethysmographically from
thoracic wall displacements is larger than that measured from integrated airflow in
patients with chronic airflow obstruction [12]. Similarly, Islam and Ulmer [27] have
shown that the plethysmographic change in apparent Rtot that occurs with airway closure
in patients with chronic airflow obstruction may become manifest in some cases at
38
WHOLE-BODY PLETHYSMOGRAPHY
volumes greater than FRCpleth. Thus the tidal volume measured from thoracic wall
displacements must also be larger than that derived from integration of airflow in some
patients with chronic airflow obstruction.
This relates importantly to the work of O’Donnell and co-workers [57–59] who have
shown that a significant limitation to exercise in patients with chronic airflow obstruction
relates to the severe dyspnoea that occurs when end-tidal inspiration encroaches on TLC.
This implies that thoracic muscle volume displacements are an important limiting factor.
The work performed by thoracic respiratory muscles includes not only flow resistive
work, but also that required to move the elastic structures of the thoracic wall itself. The
thoracic muscles move the thorax, and their volume displacements are those of the
thoracic wall, including compression and decompression of trapped gas. The volume
displacement of the thorax is systematically underestimated by integration of airflow in
patients with chronic airflow obstruction. It can be appreciated by a plethysmographic
measure of thoracic volume displacements, such as is readily available from the pressurecompensated integrated-flow whole-body plethysmograph.
Thus, it would appear that respiratory limitation in patients with chronic airflow
obstruction may be explored in the future by stimulated ventilation in an appropriate
whole-body plethysmograph. Efficacy of treatment interventions, whether pharmaceutical or rehabilitative, may be assessed by their effects on the ability of patients with
chronic airflow obstruction to improve thoracic volume displacements. Alternatively,
treatment efficacy may be assessed by changes in plethysmographic closing volume
relative to TLC in such patients.
Summary
The aim of this chapter has been to describe the unique and clinically relevant
information provided by whole-body plethysmography. Primary among this
information is the measurement of absolute TGV. Plethysmographic TGV (FRCpleth)
is considered the gold standard of absolute volume measurements and includes the
nonventilated airspace. Because the whole-body plethysmograph provides a measure
of true change in TGV, an increased use of the combination pressure-corrected
integrated-flow (transmural) plethysmograph is to be expected in the evaluation of
patients with chronic airflow obstruction. The use of thoracic volume measurements
rather than integrated mouth flow has provided more precise characterisation of
pulmonary mechanical parameters as a function of lung volume.
The clinical measurement of plethysmographic airflow resistance is also considered to
be the gold standard, and is more widely applied than either pulmonary resistance
measured invasively via oesophageal balloon or forced oscillatory resistance measured
noninvasively. It is emphasised that the plethysmographic measurement of resistance
requires two separate measurements: first, that of sRaw, and secondly, the measurement
of TGV itself. Both plethysmographic and forced oscillatory resistance are influenced
by the subject’s spontaneous breathing pattern and both require further complementary measurements to define more precisely the extent of pathophysiological
disturbances in patients with chronic airflow obstruction. Measurement of resistance
as a function of lung volume provides a useful extension of currently utilised
methodology and more clearly delineates effects of small airway obstruction.
Technological developments have now permitted incorporation of the transmural
function in commercially manufactured plethysmographs, thereby expanding the
utility of whole-body plethysmography, and increasing its utility in distinguishing
39
M.D. GOLDMAN ET AL.
between flow resistive and compression effects, both dynamic airway compression and
airway closure (nonventilated airspaces). While this capability has hitherto been
utilised primarily in FVC efforts, increased interest in new treatments for COPD may
stimulate use of this capability during tidal breathing. Whole-body plethysmography
may be further developed to include measurement of TGV during tidal breathing
without panting efforts against a closed airway shutter, and measurement of
instantaneous Raw during tidal breathing.
The sensitivity of plethysmography imposes demands for vigilance on the operator,
who must ensure stable body posture, attention to physical support of the oral cavity
and cooperation of the subject during testing procedures. Cooperation may be
improved by careful instructions to the patient, careful attention to the patient during
testing and informing the patient that they can remove the mouthpiece if breathing
becomes obstructed or too difficult. Posture must be supported to maintain subject
comfort and the instrument mouthpiece must be brought to an appropriate level for
the subject to avoid unusual neck posture. The usual clinical testing procedure of at
least three replicate measures may be usefully augmented by increased testing
replicates in circumstances where acute response to intervention is desired.
Keywords: Airway resistance, shift-volume, thoracic gas volume.
Acknowledgement. This chapter is dedicated to the life of David William Goldman, MD,
whose courage in the face of fatal lung disease inspired his father Michael David Goldman to
commit his life to efforts to make lung function testing more accurate and easier to perform
for all patients with lung disease.
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43
CHAPTER 3
Control of breathing
P.M.A. Calverley
Correspondence: P.M.A. Calverley, University Hospital Aintree, Liverpool, UK.
Most respiratory clinicians recognise that the control of breathing is a complex and
multi-factorial process that appears to be rather mysterious. Everyone would agree that
the optimum matching of ventilation and perfusion within the lungs is necessary to
maintain appropriate homeostasis for the arterial blood gases and hence for oxygen
delivery to the tissues. How this is achieved continues to elude a simple and
comprehensive explanation. For instance, the mechanisms whereby ventilation increases
appropriately and almost immediately at the start of exercise continue to be controversial
with advocates of both peripheral and central mechanisms [1]. Given this lack of a simple
understanding of ventilatory control in some of the commonest physiological situations,
it is not surprising that when disease intervenes it is hard to develop tools which are
clinically useful.
Many diseases derange blood gas tensions and these remain one of the most useful
practical ways of assessing the ability of the respiratory control system to meet the
demands placed upon it. In practice, most clinicians think about changes in blood gases
in terms of the structural processes that have produced them, rather than as
manifestations of deranged ventilatory control. Likewise, it is much easier to measure
directly, or to infer, abnormalities of lung mechanics than it is to evaluate the role that
they play in disordered control mechanisms. With a few well-recognised exceptions,
abnormal ventilatory control is not the primary reason why patients present to the
respiratory physician. Nonetheless, an understanding of the factors that influence the
control of breathing and the way in which it can modify the patients behaviour is helpful.
However, the clinical need to request the tests described briefly below are rather small.
In this chapter some of the considerations relevant to an understanding of ventilatory
control in health will be examined, followed by a review of some of the approaches that
have been used to evaluate respiratory control mechanisms. Finally, some of the diseases
where abnormal ventilatory control plays a role will be mentioned.
Ventilatory control in health
Conventionally, the neural mechanisms that regulate ventilation are considered to be a
hierarchy, with more primitive processes regulated by the higher centres [2]. There is still
disagreement about whether the output of the control system is regulated to optimise
ventilation or breathing pattern [3], although systems control approaches emphasise the
advantages of reducing total respiratory system work for any overall level of activity [4].
Inevitably, most of the data regarding the mechanisms regulating these processes have
come from animal studies with parts of the system damaged or stimulated under
anaesthesia. These are very unphysiological conditions and drawing detailed interpretations about the function of the ventilatory control system from them may be misleading.
Nonetheless, a general picture has emerged about how this system is organised.
Eur Respir Mon, 2005, 31, 44–56. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
44
CONTROL OF BREATHING
The central feature of this scheme is the respiratory pattern generator, which is
believed to lie within the medulla oblongata [5]. Here, at least three groups of neurones
receive inputs from important chemoreceptors and mechanoreceptors. Those going to
the dorsal respiratory group are predominately inspiratory but are not influenced by
stretch receptor inputs. Those in the ventral respiratory group are also thought to be
inspiratory, at least in the upper part of this complex, but activate expiratory muscles in
the lower part. Finally, neurones in the pontine respiratory group appear to cover the
period during late inspiration and early expiration and are important in terms of
respiratory rhythm generation. A basic rhythmic oscillatory pattern is developed between
these neuronal groups, which acts as a form of respiratory "pacemaker" but this is greatly
modulated by the nature and severity of the inputs from the other peripheral sensors [6].
Mechanical and chemical inputs arise from changes in chest-wall movement and/or
blood gas tensions provide the feedback, which modify the intensity and timing of the
respiratory neural outputs [7, 8]. This has a number of practical consequences:
1. If the ribcage and diaphragm are unable to produce adequate ventilation then the
mechanoreceptors in the joints and intercostals muscles will increase their output to
the respiratory centre and this leads to a change in the timing and amplitude of the
neural stimulus passing through the phrenic nerve [2].
2. If alveolar ventilation is inadequate and arterial oxygen tension (Pa,O2) falls then the
peripheral chemoreceptors at the junction of the external and common carotid
arteries increase their discharge rate [7]. One of the most consistent findings in
animal studies is the relationship between the Pa,O2 and the carotid sinus nerve
traffic and this clearly shows that these receptors are only significantly activated
once the Pa,O2 falls below 8.0 kPa [9]. However, some tonic activity does persist as
this can be abolished by oxygen therapy and this may be particularly important
during exercise.
3. Increases in arterial CO2 tension (Pa,CO2) stimulates both peripheral, but especially
central chemoreceptor, firing rate [10]. It is still difficult in humans to distinguish the
effect of CO2 on the central chemoreceptors from those on the peripheral chemoreceptors and brief stimulation, as used in many tests of ventilatory control, may not
give an appropriate index of more sustained hypercapnia, as happens in disease [11].
4. The degree of acid base compensation within the cerebrospinal fluid is an important
determinant of the effects of raised carbon dioxide and this is relevant for those
individuals who develop persistent hypercapnia or slowly worsening hypercapnia in
the course of an episode of respiratory failure [12].
5. All of the above are influenced significantly by the higher centres, as is demonstrated
by the change in ventilatory control that occurs with the onset of sleep. In these
circumstances chemical and mechanical stimulation produces less in the way of
increased ventilation and this appears to be related to the depth of sleep, with stages
three and four being those in which the patient is least responsive [13, 14]. There is
continuing debate about which part of rapid eye movement sleep is associated with
stimulation and which with depression of ventilation, although in general periods of
tonic rapid eye movement sleep has a lower responsivity than when the eye
movements are most brisk.
Thus, the integrated function of the respiratory system can be modified significantly by
changes in Pa,O2, Pa,CO2, alterations in the impedance of the respiratory system and
changes in conscious level. In health, the system appears to be regulated so as to minimise
the overall energy expenditure of the respiratory system [4]. In practice this may involve a
trade-off between the increased mechanical impedance and the disordered blood gas
45
P.M.A. CALVERLEY
tensions. One further consequence of this is the close relationship between the situation
of conflict and the sensation of breathlessness. This promotes behavioural changes,
which are often the most effective way of decreasing ventilatory demand and hence the
burden on the respiratory system [15]. For reasons of muscle energetics it is not sensible
to push the activation of the inspiratory or expiratory muscles to the point where muscle
fatigue develops and most patients either voluntarily or involuntarily adopt a breathing
pattern and a pattern of behaviour that prevents this happening [16, 17].
Assessing ventilatory control
There is no simple single method of assessing how breathing is controlled. The
investigator either opens the "closed loop" of physiological control, for example by
adding carbon dioxide to the inspired gas, or measures the breathing pattern under
resting or challenged conditions and compares this with the pattern seen in normal
subjects in similar situations. In both cases it is assumed that a brief test lasting only a few
minutes to half an hour will give information that explains how ventilation is controlled
over much longer periods. This is clearly not the case and longer exposures to abnormal
gas mixtures and mechanical loads produce different and sometimes contradictory effects
to those seen during the acute test [18]. Thus, most tests of ventilatory control give
insights into the immediate response to changes in breathing conditions but are
unreliable guides to the long-term control of breathing.
The testing of ventilatory control has gone through several phases. In the 1950s and
1960s attention focused on abnormalities in the blood gas tensions and investigators
studied the response to raised carbon dioxide and lowered oxygen in the inspired gas. In
the 1970s studies of the way in which ventilation changed in the face of added respiratory
loads similar to, or greater than, those seen in disease became an important way of
exploring control mechanisms and at this stage studies of respiratory muscle fatigue
became popular. In the 1980s longer term exposures to abnormal gas mixtures were also
investigated and highlighted the contradictory finding that increases in ventilation with
chronic hypoxia decline in intensity with more prolonged exposure, although in most
studies to date chronic hypoxia produces a level of ventilation that is greater than that at
rest. A simple approach advocated in some laboratories is the use of the breath-holding
time, although this has limitations, which will be outlined below.
With the exception of the breath-holding time, all the tests relate an external stimulus
e.g. loading (change in inspired gases) to measures that directly or indirectly reflect
respiratory centre output. The most widely used of these is minute ventilation, although
some authors have advocated reporting changes in tidal volume and respiratory
frequency separately. Unlike other tests of respiratory function it has been difficult to
establish a normal range among younger, let alone older adults. This significantly
restricts the clinical usefulness of these tests.
Measuring the output
Minute ventilation and breathing pattern
Minute ventilation can be measured directly using a Douglas bag or gas meter or can
be derived from integration of the flow signal using a flow sensor at the mouth. The
presence of a noseclip and mouthpiece influences resting ventilation and tends to increase
it possibly as a result of the added dead space [19]. This has stimulated research on
46
CONTROL OF BREATHING
TV camera
Motion
analyser
Markers
position
Geometrical
models
Flow
Volume
computing
Poes
Pga
Chest wall compartmental
volume changes
(Vrc, Vab, Vcw)
Fig. 1. – Schematic presentation of a data acquisition system used for the noninvasive measurement of lung
volume by optoelectronic plethysmography. With the subject seated with the arms supported away from the
chest light from 89 reflective markers positioned around the chest wall is collected by the television cameras and
processed to derive the volume of the space enclosed by a series of triangles constructed by the computer from
the markers position on the chest. This allows the absolute lung volume and the volume enclosed by the rib
cage and abdominal compartments of the chest wall to be calculated. Poes: oesophageal pressure; Pga: gastric
pressure; Vrc: volume of the rib cage; Vab: volume of the abdomen; Vcw: volume of the chest wall.
noninvasive methods of measuring ventilation and for a period the magnetometer
approach of quantifying the excursion of the ribcage and the abdomen became popular
[20]. Commercial devices incorporating impedance coils into bands placed around the
chest and abdomen were used to simplify the application of the system. Unfortunately, it
proved difficult to calibrate this system and changes in position of even a minor degree
significantly affected the reliability of the measurement. More accurate noninvasive
measurements are now available using optoelectronic plethysmography [21, 22]. To date
these have only been used in a research setting (fig. 1).
Analysis of breathing patterns
Breathing patterns can be characterised simply from the combination of tidal volume
and respiratory frequency but can be further separated into mean inspiratory flow (tidal
volume divided by inspiratory time) and the duty cycle (tidal volume divided by total
cycle duration) [23]. The characteristics can also be used to define minute ventilation. In
general, the mean inspiratory flow is the variable that responds to external stimulation,
although when mechanically limited this can decline in the face of a rising neural drive. In
this context it can be more useful to measure the mouth occlusion pressure (the pressure
recorded 100 ms after the onset of inspiration against a closed airway). In this brief
period the muscles should shorten simply in response to the pattern of impulses in the
phrenic nerve [24]. Unfortunately, technical considerations related to the shape of the
diaphragm and to identifying reliably the onset of the occlusion of pressure waveform
make this technique less robust in disease than it is in healthy subjects. Nonetheless, it is a
significant improvement on simply reporting tidal volume in patients with abnormal
47
P.M.A. CALVERLEY
resting lung mechanics. An alternative tactic that is less well validated is to express the
tidal volume or minute ventilation as a per cent of that predicted [25]. Formulae exist for
deriving the predicted maximal minute ventilation and some data suggest that this
approach may be useful [26]. However, the lack of widespread normal values limits its
application.
A further, more direct but unfortunately more invasive, method involves the use of a
balloon catheter system placed in the oesophagus and stomach from which the
transdiaphragmatic pressure can be calculated. Even if this is simplified to just an
oesophageal balloon it is an uncomfortable procedure that is rather time consuming. The
advantage is that this should indicate the point at which inspiration begins and from
which the occlusion pressure waveform can be calculated. Unfortunately, this has
problems in patients with chronic obstructive pulmonary disease (COPD) where there is
a degree of intrinsic positive end-expiratory pressure, which makes it hard to be sure at
which point the neural inspiration has begun [27].
A yet more invasive approach is to monitor diaphragmatic electromyogram, which
gives a measure of the electrical activation [28]. Unfortunately, relating the size of the
electrical signal to either the pressure developed or the mechanical change it produces can
be very difficult [28].
Chemosensitivity responses to hypoxia and hypercapnia
As noted previously, the arterial blood gas tensions at rest often provide the simplest
indicator of the adequacy of ventilatory control. The role of abnormal gas exchange is
considered elsewhere in this Monograph. Nonetheless, in many circumstances,
particularly those where there is little or no mechanical abnormality, an elevation in
the carbon dioxide tension is a good pointer of inadequate ventilation and impaired
control mechanisms.
Protocols for stimulation either by increasing carbon dioxide or reducing the inspired
oxygen concentration normally involve the subject rebreathing from a sealed anaesthetic
bag, which containsy6 L of premixed gas, and then recording the change in gas tensions
at the mouth and relating it to either ventilation or occlusion pressure [29, 30].
Fortunately, the relationships between a fall in oxygen saturation and ventilation are
themselves linear and hence peripheral oxygen saturation can be measured as a useful
surrogate for Pa,O2. Linearised in this way makes the slope of the ventilation saturation
relationship, which is an inverse one, as useful a guide to chemosensitivity as is the slope
of the ventilation–CO2 relationship [30]. In health, wide ranges exist for both values and
it is important to specify the hypoxic response in terms of the CO2 tension maintained
during the trial. Isocapnic testing is essential in studies of hypoxic response and is
normally achieved by re-circulating the expired gas through a carbon dioxide scrubber
circuit [30]. Likewise, oxygen tensions during the CO2 re-breathing are an important
practical consideration and patients normally begin with a gas mixture containing 93%
oxygen and 7% CO2.
Mechanical loading
This is a more specialised and largely research technology, which might yet prove
useful if developed as a clinical tool. A variety of loaded breathing circuits have been
developed involving either exposure to a single load and observing the effect on the first
loaded breath relative to the preceding breaths or studying the effects of stimulated
48
CONTROL OF BREATHING
breathing with hypoxia or hypercapnia during a fixed resistive load [25]. It is more
difficult to apply sustained elastic loads and so most data derives from studies during
resistive loading. Although it would be intellectually interesting to understand the effects
of stiffening of the chest wall, such studies would be unlikely to produce different data
from that during resistive breathing. Again, the nature of the disease to be studied affects
the results obtained and in this context ventilatory restriction due to mechanical
impairment of the chest wall or the effect of intrinsic positive end expiratory pressure in
COPD can complicate the interpretation of these tests [31]. When resistive loading is used
it is important to know the linearity of the resistance as higher levels of ventilation may
expose the patient to substantially greater inspiratory loads than is initially anticipated.
While minute ventilation is normally depressed in the face of inspiratory resistive
loading the degree that subjects defend their ventilation varies significantly [32].
Likewise, individuals will adopt rather different breathing patterns when acutely loaded.
The relevance of the breathing pattern adopted to more chronic loading is less clear and
in general loaded breathing produces a reduction in tidal volume and a shortening of
respiratory cycle duration i.e. a rapid, shallow breathing pattern [33–34]. This seems to be
true whether the load is predominantly resistive or elastic. Increasing knowledge of the
abnormalities of lung mechanics in many conditions suggests that in life the loaded
breathing is a combination of the two.
Breath-holding time
This simple test requires only a noseclip and a stop-watch. The subject exhales to
residual volume, inhales to total lung capacity and then holds the breath for as long as
possible. Analysis of the expired gas allows the stimulus to changing gas tensions to be
related to the breath-holding time. The test can be repeated until reproducible results are
obtained or subject exhaustion intervenes [34]. Like many other tests of ventilatory
control there are a number of other important confounders that interfere with the
interpretation of this test. Specifically, respiratory muscle strength and the geometry of
the chest wall are very relevant to the patient’s ability to hold the breath [35, 36]. If this
test is to be used clinically it is necessary that the laboratory establish its own normal
range so that some comment on when the breath-holding time is abnormal can be made.
Ventilatory control and disease
Problems of interpretation
A variety of factors interfere with the interpretation of ventilatory control and these
are well illustrated from studies in COPD, which is one of the most examined conditions.
In general, these can be summarised as:
1. Impaired gas mixing: in COPD and chronic asthma, lung gas stores are increased
and mixing of fresh and resident gases is not homogeneous. Slowly ventilated areas
of the lungs may not achieve the same gas concentrations as the well ventilated ones,
with a discrepancy resulting between measurements made on end tidal expiratory
breaths and the true gas tensions within the alveoli.
2. Mechanical inequalities: abnormal time constants throughout the respiratory system
(the product of resistance and compliance) can delay the equilibration of pressure
applied within the chest and that recorded at the mouth. This is a particular problem
49
P.M.A. CALVERLEY
for obstructive lung disease, but is not an issue for interstitial lung disease where the
time constants are short.
3. Changes in the respiratory muscles: changes in the configuration of the respiratory
muscles can greatly influence their ability to develop pressure for a given neural
stimulus [37]. Thus, patients with COPD who have a flattened diaphragm, which
shortens this muscle, will be unable to develop the same pressure for an equivalent
stimulus as would someone with interstitial lung disease where the diaphragm has a
normal configuration. This is very relevant when indices, like mouth occlusion
pressure, are used as the surrogate for respiratory centre output. Moreover, changes
in muscle structure occur in chronically shorted muscle, like the flattened diaphragm
in patients where lung volume is increased [38], although how relevant these changes
are to human disease is uncertain.
4. Mechanical loading by preventing muscle shortening can itself reduce outputs such
as minute ventilation, irrespective of the stimulus being applied to the respiratory
muscles. This problem has been noted previously.
Specific diseases
Primary disorders of ventilatory control
The assessment of ventilatory control can be a useful adjunct to the investigation of
patients in a number of clinical settings. The two most important of these are in patients
who either under breathe (hypoventilate) or over breathe (hyperventilate). In each case
the investigation has been mainly used as a research tool rather than one required to
make a clinical diagnosis. In the relatively rare primary alveolar hypoventilation
syndrome, which is seen mainly in children, there is an absence of chemoreceptor
responsiveness [39]. This is compensated for during the day by the wakefulness-related
drive to breathe, as commented on previously. But at night when this influence declines,
so does the minute ventilation with profound falls in Pa,O2 and an increase in Pa,CO2.
These subjects show a reduced response to CO2 and hypoxic stimuli even when awake,
and this becomes much more dramatic when they are asleep. Why this condition arises is
still problematic, but it commonly presents in childhood or early adolescence with failure
to thrive, intellectual impairment, daytime tiredness or even light heart failure and
pulmonary hypertension. Appropriate therapy with nocturnal ventilation can dramatically improve the well being of these patients and permits adequate correction of their
blood gas tensions [40].
The problems of disproportionate breathlessness associated with individuals who
appear to develop a larger ventilatory response than required during exercise, have been
known for many years. More recently, careful scientific study has shown that these
patients have abnormal ventilatory control [41]. They can usually be identified either by a
low Pa,CO2 and a high Pa,O2 at rest and will commonly have larger tidal volumes and
higher respiratory frequencies than would be predicted during progressive exercise
testing. They exhibit ventilatory limitation in this setting, and some who are not
hypocapnic at rest will become so during the exercise, although this tends to resolve as
their mechanical inability to sustain high ventilations eventually limits their capacity to
continue. Why these patients behave in this way is unclear, but their ability to sustain the
hyperventilation for extended periods during the daytime and overnight suggest that
there is a primary problem with ventilatory control, rather than simply a secondary
psychological abnormality. Undoubtedly, the chronicity of their problems produces
psychological difficulties and it has been difficult to disentangle cause and effect in these
circumstances.
50
CONTROL OF BREATHING
A third situation when ventilation will be abnormal arises with periodic breathing.
This is now known to be a relatively frequent finding during sleep in patients with chronic
congestive heart failure, having been originally recognised only during the daytime in
patients with gross congestive heart failure, or major strokes. The waxing and waning of
ventilation with the system "hunting" to find a stable CO2 tension is characteristic of this
form of breathing. Changes in CO2 threshold, which occurred during sleep, appear to be
important for initiating this form of respiration, although the arousal responses that
commonly accompany it are also relevant [42]. At present, periodic breathing appears to
be a marker of other diseases, rather than a primary abnormality of ventilatory control,
which itself produces ill health.
Chronic obstructive pulmonary disease
This is the most widely studied condition from a research perspective and is reviewed in
detail elsewhere [43]. Attempts to distinguish patients who developed hypoxia and
hypercapnoea from those who maintained their CO2 tensions have tracked the history of
this research methodology. The initial observation that suggested that those with
hypoxia had lower ventilatory responses has been hard to substantiate. In general, the
worse the mechanical impairment in COPD the lower the ventilatory response to altered
gas tensions. The same mechanical abnormalities limit the interpretation of occlusion
pressure responses to either mechanical or chemical loading [44, 45]. The most consistent
findings have been in the breathing pattern, with patients who develop carbon dioxide
retention also having a smaller tidal volume and reduced inspiratory time [46].
Effectively, the smaller tidal volume in the face of the fixed dead space increases the dead
space tidal volume ratio and promotes carbon dioxide retention as predicted from steady
state analysis of gas exchange. Researchers continue to try and disentangle the cause and
effect nature of both respiratory sensory abnormalities and the mechanical impairment
that characterises these patients. Nonetheless, many of us who have studied these
problems are left with the feeling that there is a difference in the basic responsiveness of
patients who will readily allow themselves to become hypoxaemic and hypercapnic [47]
and this is an issue that still awaits adequate resolution.
A further controversy in the field of ventilatory control on COPD has been the effect of
high inspired oxygen concentrations. This issue has been reviewed on several occasions
[48, 49] and it is clear that some of the conflicting data reflects the severity of the disease
in which the measurements have been made. In very severe COPD, with marked
hypoxaemia and possibly associated haemodynamic instability, exposure to a high
concentration of oxygen produces CO2 retention largely by a predictable effect on
ventilation-perfusion matching within the lung [50]. In contrast, in nonventilated patients
breathing high oxygen concentrations during an exacerbation, a proportion will develop
CO2 retention because of a degree of ventilatory depression, perhaps secondary to a
reduction in chemoreceptor activation. The possibility that there has also been an effect
of the change of gas tensions on airway calibre must also be considered [51]. In practice, it
is clinically desirable to avoid this and this is still best done with controlled oxygen
therapy as described by Campbell over 40 years ago.
Bronchial asthma
No specific ventilatory control abnormalities have been identified in most patients with
bronchial asthma. However, it is clear that some asthmatics deteriorate acutely and there
is increasing data that these patients perceive changes in chemical stimuli relatively
poorly [52]. This occurs despite the fact that some appear to have reasonably normal lung
51
P.M.A. CALVERLEY
mechanics between attacks and suggest that there is either an inherent or acquired defect
in ventilatory control, which when combined with the onset of severe asthma exposes
these patients to particular risk. Early studies showed that the ventilatory response to
CO2 was reduced in some asthmatics during the recovery from a severe exacerbation [53]
and where there is chronic loading both respiratory perception and ventilatory responses
to chemical stimuli were impaired [54].
Interstitial lung disease
This heterogeneous group of conditions is associated with the increasing elastic work
of breathing and as noted previously, the breathing pattern in response to this tends to be
rapid and shallow. Arterial blood gas tensions are usually normal at rest in these
conditions but there is marked desaturation during exercise as ventilation fails to match
the increased perfusion. Typically, occlusion pressure responses are increased in these
patients, but it is unlikely that they are mounting a greater ventilatory response for a
given mechanical load than other subjects.
Chest wall and neuromuscular diseases
During the 1970s careful examination of well-characterised patients with kyphoscoliosis showed that the reduction in response to CO2, which would be anticipated in these
patients, was related to the reduction of the compliance of the respiratory system and
particularly the change in the chest wall compliance [55]. Again, the breathing pattern
was rapid and shallow, as is seen in patients with other disorders associated with
impaired lung mechanics. The management of these patients has been revolutionised by
the introduction of nocturnal positive pressure ventilation treatment but good studies
examining ventilatory responsiveness before and after the introduction of this therapy
are lacking. There is a continuing suspicion that alterations in chest wall mechanics may
explain why Pa,CO2 falls during the day, although a resetting of the chemoreceptors as a
result of better nocturnal ventilation is also a possibility. Patients with neuromuscular
disease normally show similar abnormalities in their response to CO2 and oxygen tension
perturbation as do other subjects who are unable to mount an adequate ventilatory
response [56]. These patients are particularly dependent upon the diaphragm function,
which decreases during rapid eye movement and often the earliest sign of future
ventilatory impairment is the development of oxygen desaturation in this sleep stage.
However, their overall respiratory function as assessed by their vital capacity is a better
predictor of prognosis than the presence of this isolated ventilatory control problem [57,
58].
Sleep and breathing disorders
Most attention has focused on the anatomical abnormalities, which determine the
occurrence of upper airway obstruction during sleep, rather than changes in ventilatory
control. Most clinicians accept that there must be some individual variation in response,
and alterations in ventilatory control may explain some of the variation in the severity of
nocturnal oxygen desaturation and duration of the apnoeic periods [59]. As yet, no
simple way of assessing this has been developed. The large number of complex and
changing variables, which characterise upper airway obstruction during sleep, make the
simple elucidation of this problem unlikely in the near future.
52
CONTROL OF BREATHING
Conclusion
The regulation of ventilatory control is abnormal in some relatively rare disease states
but its major role in most cases is to modify the clinical presentation and subsequent
progress of patients with many different forms of respiratory disease. Although major
steps forward have been made in understanding some of the complex interactions
between altered blood gas tensions, lung mechanics and the central nervous system
processing of these signals, it is difficult to turn this information into tools that modify
clinical decision making. It seems likely that quite different approaches to understand
ventilatory control, perhaps coming from areas of systems control theory, will ultimately
give better ways of explaining what is happening. Until such "high tech" solutions are
available, an awareness of the additional impact of altered ventilatory control is helpful
when considering patient management.
Summary
The maintenance of blood gas homeostasis is dependent on the balance between
respiratory drive and peripheral, mechanical and chemoreceptor responses. No single
measurement encapsulates all aspects of this complex control system. Most
investigators and clinical tests rely on relatively short-term changes in inspired gas
concentrations and/or additional predominantly inspiratory mechanical loading to
determine how the control system responds. Usually ventilation or an index of neural
drive, such as mouth occlusion pressure, is used as the output measurement. Changes
in the mechanical properties of the lungs make interpretation of these tests difficult
and in common diseases such as chronic obstructive pulmonary disease, asthma and
interstitial lung diseases the usual index of ventilatory control abnormality is a change
in the arterial blood gas tension. In some conditions, e.g. hypo- or hyperventilation
syndromes, investigation of respiratory control mechanisms may be useful. Studies of
disordered respiratory control have helped understanding of the pathophysiology of
disease and continue to inform current clinical practice, e.g. in the prescription of highflow oxygen. Future developments using modern computerised methods to analyse
breathing pattern and relate this to neural activation may offer more appropriate
clinical tools.
Keywords: Chemoreceptor, chronic obstructive pulmonary disease, hypercapnoea,
hypoxia, sleep disorders.
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56
CHAPTER 4
Respiratory muscle assessment
T. Troosters*,#,}, R. Gosselink*,#, M. Decramer*,#
*Respiratory Division and Respiratory Rehabilitation, Respiratory Muscle Research Unit, and #Faculty of
Kinesiology and Rehabilitation Sciences, Dept Rehabilitation Sciences, Katholieke Universiteit Leuven,
Leuven, Belgium. }Postdoctoral fellow of the Fonds voor Wetenschappelijk Onderzoek-Vlaanderen.
Correspondence: T. Troosters, Respiratory Division and Respiratory Rehabilitation, University Hospital
Gasthuisberg, Herestraat 49, B3000 Leuven, Belgium.
Support statement
This work has been supported by grants: Fonds voor Wetenschappelijk Onderzoek Vlaanderen, Grant
# G.0237.01 and # G.0175.99, and "Levenslijn" grant # 7.0007.00.
Respiratory muscles generate the pressure differences driving ventilation. Respiratory
muscle weakness is hence an important clinical feature. In advanced stages, respiratory
muscle weakness leads to respiratory pump failure. Respiratory muscle dysfunction (i.e.
reduced strength or endurance) is to be distinguished from lung function abnormalities,
and should be measured separately. Inspiratory muscle weakness may partially explain
dyspnoea and exercise intolerance. In addition, reduced respiratory muscle force has
been shown to be an important predictive factor for poor survival in chronic obstructive
pulmonary disease (COPD) [1], cystic fibrosis [2] and congestive heart failure [3]. In
advanced stages the functional consequence of respiratory muscle weakness is a reduction of the operational lung volume and patients may require mechanical ventilation.
Expiratory muscle weakness leads to problems with speech, and mucus retention due to
impaired cough efficacy.
Measurement of respiratory muscle function is important in the diagnosis of respiratory
muscle disease [4–6], or respiratory muscle dysfunction [7]. It may also be helpful in the
assessment of the impact of chronic diseases [8–12] or their treatment [13–15] on the
respiratory muscles. For example, specific inspiratory muscle training has been reported to
be useful in COPD only when patients present with significant respiratory muscle weakness [15], and tapering of oral corticosteroid treatment successfully restored respiratory
muscle strength and dyspnoea in patients with corticosteroid-induced myopathy [16].
The present chapter aims to provide clinicians with some aspects of respiratory muscle
testing. More detailed, excellent reviews on the pathophysiology and aetiology of
respiratory muscle weakness are available elsewhere for the interested reader [17, 18].
Indications, techniques commonly used in clinical practice and issues important in the
interpretation of the test results are the main focus of this chapter.
When should respiratory muscle function be assessed?
Measurements of respiratory muscle function should be performed as part of a more
complete diagnostic process including anamnesis and physical examination, arterial
blood gas analysis and imaging techniques. Lung function assessment including
spirometry, assessment of static lung volumes, and diffusion capacity further completes the technical investigations relevant in the diagnostic process. Measurements of
Eur Respir Mon, 2005, 31, 57–71. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
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T. TROOSTERS ET AL.
respiratory muscle strength or endurance should never be over-interpreted. A low
inspiratory or expiratory muscle strength without clinical context has relatively poorly
defined clinical consequences and the range of normality in healthy subjects is very large
[19]. The clinician may encounter two possibilities that would prompt for careful
assessment of respiratory muscle function: 1) clinical signs or symptoms that are
suggestive of respiratory muscle weakness; or 2) a pathological condition where
respiratory muscle weakness may occur and assessment of the respiratory muscles is
advised in the screening, prevention, or follow-up of these patients.
Clinical signs of respiratory muscle weakness
Clinical signs or symptoms that can be suggestive of respiratory muscle weakness can
be summarised as follows: 1) unexplained reduction in vital capacity; 2) CO2 retention
while awake or during sleep, specifically in the absence of severe airflow obstruction; 3)
shortness of breath; 4) orthopnoea (shortness of breath while supine), or dyspnoea
during bathing or swimming; 5) short sentences during speech; 6) tachypnoea; 7)
paradoxical movement of the abdominal or thoracic wall; 8) problems with cough (and
recurrent infections); and 9) generalised muscle weakness.
Respiratory muscle weakness is often advanced before clinical symptoms occur. This
follows from the relatively low respiratory muscle force that is required to overcome
most respiratory tasks. In addition, symptoms only poorly relate to measurements of
respiratory muscle strength or endurance. In patients with neuromuscular disease, for
instance, hypercapnia only modestly relates to respiratory muscle strength [5, 20]. This is
due to the fact that symptoms generally only occur in the presence of an imbalance
between the load on the respiratory pump and its capacity [21]. Respiratory muscle
function measurements address only the latter.
When respiratory muscle strength is moderately to severely reduced, discrete clinical
symptoms may occur, and this may prompt for assessment of the respiratory muscles to
help in the diagnostic process. The cardinal symptom of respiratory muscle weakness is
dyspnoea. When muscle weakness becomes more obvious, symptoms may also occur at
rest, dyspnoea, hypercapnia and/or speech problems disable the patient. In the case of
severe expiratory muscle weakness, reduced cough efficiency may become an important
handicap and patients may become ventilator dependent. Only in severe respiratory
muscle dysfunction, vital capacity is generally reduced as a consequence of the
respiratory muscle weakness and may become a better predictor of morbidity than
measurements of respiratory muscle strength [22].
Pathological conditions in which respiratory muscle weakness can be suspected
Patients with neuromuscular or metabolic diseases are obviously at risk to develop
skeletal and respiratory muscle weakness. In some cases the respiratory muscle weakness
and related symptoms are even the first presenting symptoms [23, 24]. In neuromuscular
diseases close attention should be paid to the involvement of both the inspiratory and the
expiratory muscles. In patients with multiple sclerosis for example, abdominal (and hence
expiratory) muscle weakness is a hallmark of the disease [25], and is related to clinical
problems, such as mucus retention. In lung diseases, such as cystic fibrosis and COPD,
inspiratory muscle weakness is often present [26]. As a contradiction at first sight,
respiratory muscles seem, on average, to be relatively well trained in these diseases [27–
29]. The low respiratory pressures are due to the mechanical constraints and
hyperinflation rather than to pure muscle weakness. When patients are malnourished
or exposed to corticosteroids, however, weakness of the respiratory muscles is seen in
these diseases [13, 30, 31]. Some attention has recently been given to expiratory muscle
58
RESPIRATORY MUSCLE ASSESSMENT
weakness in obstructive lung diseases. Abdominal muscle strength was more than normal
in cystic fibrosis [27], probably as a consequence of the chronic coughing. In COPD
patients, expiratory muscle weakness is seen frequently [32, 33], but the clinical
importance of it is not well understood.
Less obvious, but nonetheless important is the detection of respiratory muscle
weakness in patients with heart failure [9], cancer [34] and systemic diseases, such as
sarcoidosis [12, 35, 36]. In patients diagnosed with hyperventilation [37] and asthma [38,
39], respiratory muscle weakness can contribute to the sensation of dyspnoea and the
assessment of respiratory muscle function may be helpful in solving the diagnostic
dilemma of unexplained dyspnoea.
When patients are treated with drugs that may induce myopathy, it may be prudent to
assess respiratory muscle strength before initiating the treatment, and proper follow-up
of patients is advised [40]. After corticosteroid treatment, respiratory muscle function is
often impaired [16], and long-term colchicine treatment may also induce respiratory
muscle weakness [41]. Hence the assessment of respiratory muscle function is surely not
restricted to patients with pulmonary diseases.
Principles of assessment of respiratory muscles
Measurement of respiratory muscle strength is no novelty in the lung function
laboratory [42]. It is nowadays routinely performed in clinical practice as tools have
become available that allow these measurements to be performed routinely in clinical
practice. The interpretation of measurements of respiratory muscle strength, however,
may be somewhat more complex than most other measurements of skeletal muscle
strength, for the reasons described below.
In clinical practice respiratory muscle force is indirectly measured through the pressure
generated during inspiration or expiration. Respiratory muscle force is generally
expressed as kilopascal (kPa) or cm water pressure (cmH2O: 1 kPa=10.2 cmH2O). These
pressures reflect pressure changes against atmospheric pressure. The pressure is
generated by all the muscles under investigation (inspiratory or expiratory), and is
hence not muscle specific. In addition, reduced respiratory muscle force may result from
cerebral, spinal cord, anterior horn, peripheral (i.e. phrenic) nerve, neuromuscular
junction or the muscle fibre dysfunction. At each level pathology may occur and hence
reduced respiratory pressures should not be necessarily attributed to a respiratory muscle
dysfunction per se. The pressures measured depend also on the geometry of the thorax in
which the pressure is generated. For instance, the pressure generated by the diaphragm is
dependent of its in vivo three dimensional shape taking into account: 1) Laplace law (in
brief, the Laplace law implies an inverse relationship between the radius and the
pressure); 2) the relative degree to which it is apposed to the rib cage; and 3) its length
force properties [43]. In stable patients with emphysema, the "flattened" diaphragm often
fails to generate normal pressure, although the diaphragm muscle is generally believed to
be well "trained" [28, 44–46].
Another variable influencing the outcome of the inspiratory and expiratory pressure measurement is the relative lung volume at which it is obtained. Like all skeletal
muscles, the respiratory muscles have a well defined length–tension relationship. If the
diaphragm is shortened below its optimal length (L0, the length at which a maximal
tension is obtained) it can generate less tension [47]. This has repercussions during
acute hyperinflation, where the mechanism of reduced tension generating capacity
of the diaphragm seems to be more important than the geometric changes [48]. The
length–tension relationship has important consequences for the technique of measuring
59
T. TROOSTERS ET AL.
in- and expiratory muscle force. Indeed, changes in the lung volume at which the
measurement is performed may alter the outcome of the measurement. Hence, lung
volumes should be properly standardised.
A final factor, related to the above, influencing the pressure measured during maximal
inspiratory or expiratory manoeuvres, is the elastic recoil of the lungs (inward) and chest
wall (outward). At the functional residual capacity the elastic recoil of the lungs and the
chest wall generate equal pressures. Hence, at this lung volume any additional pressure
measured during in- or expiration originates exclusively from respiratory muscle
activation. When expiratory pressures are measured at total lung capacity (TLC), the
recorded pressures are the result of the expiratory muscle and the elastic lung recoil at
TLC. Alternatively, when inspiratory pressures are assessed at residual volume (RV), the
resultant pressures originate from the action of the inspiratory muscles, and the pressure
generated by the tendency of the chest wall to expand at RV.
Taking the above into account, clinicians should be aware that the respiratory
pressures obtained in patients, or healthy subjects are not a "clean" measure of the
strength of the respiratory muscles. They are the net result of the tension (force)
generated by the muscle, which is dependent on the lung volume at which the manoeuvre
is obtained and the chest wall and lung mechanics. Elastic recoil is also dependent on the
lung volume, but may also be altered by the disease (e.g. lung fibrosis versus emphysema).
The resulting pressures are, however, a good reflection of the functional reserve of the
respiratory pump, since the net pressure generated is needed to drive the ventilation.
Measuring respiratory muscle force
Measurements of respiratory muscle function are generally obtained from measuring
pressures achieved by volitional activation or electrical or magnetical stimulation of the
phrenic nerve or motor roots. Pressure can be measured in the nose, at the mouth, in the
oesophagus, or across the diaphragm (measuring the pressure above, in the oesophagus,
and below the diaphragm, in the stomach). Lung function impairment (static and
dynamic lung volumes) does not correlate with respiratory muscle dysfunction, with the
exception of patients with neuromuscular disease in advanced stages. Techniques used in
the lung function laboratory are described below.
Maximal voluntary respiratory pressures measured at the mouth
Maximal voluntary inspiratory (PI,max) and expiratory (PE,max) pressures (or MIP and
MEP) are probably the most frequently reported noninvasive estimates of respiratory
muscle force. Ever since Black and Hyatt [42] reported this noninvasive technique in
the late 1960s it has been widely used in patients, healthy control subjects across all ages,
and athletes. Pressure is recorded at the mouth during a quasi-static short (few seconds)
maximal inspiration (Müller manoeuvre) or expiration (Valsalva manoeuvre). No airflow
is allowed during the manoeuvre and pressure can build up to w30 kPa in extremely fit
healthy subjects. The manoeuvre is generally performed at RV for PI,max, and at TLC for
PE,max. Although functional residual capacity would theoretically be more appropriate,
as lung and chest wall compliance are neutralised, and the pressure theoretically would
better reflect the tension developed by the respiratory muscles (Pmus), patients find it
easier and more straightforward to perform the manoeuvres from RV and TLC. Only
few contraindications exist for these measurements and these can be summarised as
pathological conditions where relatively large pressure swings in the thorax or abdomen
60
RESPIRATORY MUSCLE ASSESSMENT
should be avoided (e.g. aneurism, uncontrolled hypertension, urinary incontinence). The
coefficient of variation is reported to be acceptable for a clinical test (6–9%) [49–51].
Although the technique appears simple at first sight and hard- and software became
available to make these measurements easily accessible in the pulmonary function
laboratory, there are some technical pitfalls that may influence the obtained results and
make the results more variable than most other lung function measurements. Some
critical aspects in the methodology are summarised below.
Tracing inspection. Quality control of the measurements can only be obtained from
inspection of the pressure–time curves. The peak pressure should be obtained in the very
beginning of the manoeuvre. The pressure maintained for at least 1 s is generally reported
as the PI,max or PE,max (plateau pressure) [17]. A recent study, however, challenged the use
of the plateau pressure, concluding that the peak pressure may be easier to obtain and
equally reliable when subjects are well instructed [52].
Position. Measurements are obtained preferably in the sitting position. Although body
posture has no significant influence on the result of the measurement in healthy subjects
[53], and even in convalescent neonates [54], in COPD patients changes in body posture
may significantly impact on the obtained result. Leaning forward for example may result
in higher inspiratory pressures [55], while measurements obtained in the recumbent
position may lead to lower pressures [56].
Leak. To avoid pressure generation by the muscles of the cheeks and buccal muscles, a
small leak should be present in the equipment. The leak described by Black is 15 mm long
and has an internal diameter of 2 mm. Using this leak, the glottis should be opened to
generate pressures for w1 s, and the pressure obtained reflects the pressure generated by
the respiratory muscles. When a leak is absent, the recorded pressures may erroneously
reflect the pressure generated in the mouth by the cheeks and buccal muscles.
Mouthpiece. Flanged mouthpieces (as the ones generally used for lung function testing)
have been reported to result in pressures inferior to those obtained when a rigid
mouthpiece is sealed against the mouth. Especially for expiratory pressures, flanged
mouthpieces may result in underestimated pressures due to additional leaks that appear
with the increased pressure in the mouth [57]. Sometimes tests can be more successfully
performed using a face mask (especially in patients with neuromuscular diseases
characterised by facial or bulbar muscle weakness). On average there is no significant
difference in PI,max, but PE,max may be higher using a tube or nonflanged mouthpiece [58].
Practice tests. Tests should be performed by an experienced technician. Since the
Valsalva or Müller manoeuvres are unfamiliar to patients the manoeuvres should be
carefully explained. There has been debate on the number of repetitions that need to
be carried out before a result can be considered valid [59–62]. The current authors’
experience, shared by others [19], suggests that a minimum of five manoeuvres should be
performed, and reproducibility should be within 5–10%. Increasing the number of
measurements is time consuming and tedious. In case of questionable effort, a sniff nasal
pressure manoeuvre (see below) may give additional information.
Equipment. A recent statement of the American Thoracic Society and European
Respiratory Society advises to use metal membrane or piezoelectric transducers with an
accuracy of 0.049 kPa (0.5cmH20) in a pressure range of ¡19.6 kPa (¡200 cmH20). When
healthy subjects are tested, higher expiratory pressures may be obtained. In a cohort of 85
61
T. TROOSTERS ET AL.
healthy subjects, tested in the current authors’ laboratory and agedw50 yrs, the maximum
inspiratory and expiratory pressure obtained were -17.6 kPa (-180 cmH2O) and 30.2 kPa
(308 cmH2O) respectively.
It is preferred that the signal of pressure versus time is recorded, and is available to the
technician for immediate inspection. Calibration of the manometer should be carried out
regularly, and can be done easily using a water column. Mercury is preferably not used
due to contamination problems.
Interpretation and normal values. In absolute numbers, the PE,max is roughly the double
of PI,max when the Black and Hyatt technique is used, with a rigid mouthpiece. In this case it
is very rare to find PE,max inferior to PI,max. This is illustrated in figure 1. However, in some
diseases (e.g. spinal cord injury, below C3-5, multiple sclerosis) PE,max is typically more
reduced than PI,max, and the value of PE,max may be inferior to PI,max (fig. 1). In addition,
when a flanged mouthpiece is used, the PE,max may often be underestimated due to leaks.
Many authors have reported normal values for PI,max and PE,max. Impressive
differences are observed between the normal values [19, 58, 62–71] reported in the
literature. This variability is depicted in figure 2 where an overview of available sets of
normal male subjects is given as a function of age. Roughly, it can be seen that there
is a decline of inspiratory muscle force from the age of 20–25 yrs. Hence if children are
tested, separate normal values are advised. This is largely due to the previously described
differences in methodology (lung volume, mouthpiece, number of repetitions). It is
advised that a cohort of healthy subjects is tested and consequently the most appropriate
35
l
30
l
PE,max kPa
l
l
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l l
ll l
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l l
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l
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l l l
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lll l
lll
l l
lll
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l
ll
ll
l
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l l
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ll l
l ll
l
ll l l
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25
l
20
15
10
5
0
l
ll
0
5
10
15
PI,max kPa
20
25
30
35
Fig. 1. – Maximum inspiratory and expiratory pressure (PI,max and PE,max) measured in 85 healthy subjects (#),
21 patients with multiple sclerosis (MS; $) tested in the current authors’ centre [99], and 13 patients with spinal
cord injury (SCI; h) [100]. As can be observed, in healthy subjects the PE,max exceeds the PI,max in every single
case. In MS, PI,max may be larger than PE,max, and in SCI, PI,max is typically larger than PE,max.
62
RESPIRATORY MUSCLE ASSESSMENT
20
s
s
l
u
l
n
u
s
X
s
s
l
s
10
l
s
l l
s
PI,max kPa
15
u
u
n
s
u
5
0
0
10
20
30
40
50 60
Age yrs
70
80
90
100
Fig. 2. – Predicted normal inspiratory pressures measured at the mouth for healthy male subjects as reported
from the different cohorts reported in the literature. Maximum inspiratory preasure (PI,max) is reported in
cmH2O, age in years. Symbols represent different studies: %: Wijkstra et al., 1995 [62]; ,: Uldry and
Fitting, 1995 [72]; ': rochester and Arora, 1983 [64]; &: Hautmann et al., 2000 [70]; $: Heijdra et al.,
1994 [56]; h: Enright et al., 1994 [19]; (: Vincken et al., 1987 [65]; #: Leech et al., 1983 [68]; ): Wilson
et al., 1984 [67]; 6: McElvaney et al., 1989 [69]; z: Ringqvist, 1966 [66].
reference values are chosen. In addition, it has to be noted that in all models of maximal
in- and expiratory pressures the explained variance is low, reflecting large inter-individual
differences even when age, sex and anthropometric values are taken into account. Hence,
a low PI,max should always be interpreted with caution. A normal PI,max, however,
generally excludes clinically relevant inspiratory muscle pathology.
Inspiratory pressure measured at the nose
PI,max measured at the nostril Psniff during a sniff manoeuvre is a relatively newly
developed technique [73] to measure inspiratory muscle function. One of the main
advantages is that it is a technique that involves a natural manoeuvre (sniff), which is
"easy to understand" by the patient [74]. Pressure is measured in an occluded nostril
during a forced sniff. The unoccluded nostril serves as a variable resistance, prohibiting
flow w30 L?min-1, and the pressures measured at the nose reflect those obtained in the
oesophagus during sniff manoeuvre [73]. Since there is more airflow compared with the
PI,max manoeuvre, these sniff manoeuvres are not static. Generally the sniff nasal
pressures are as high as PI,max (or even slightly higher) [72]. Maillard et al. [49] reported
a Psniff/PI,max ratio of 1.03¡0.17, and reported equal and good within session
reproducibility. Although less common in routine clinical practice this technique
showed to be extremely useful in the diagnosis and follow-up of respiratory muscle
weakness in children [75, 76], and patients with neuromuscular disease [77, 78] where
sniff nasal pressures were reported to be superior to PI,max. It should be acknowledged
that some investigators reported sniff nasal pressures to be inferior to PI,max in severe
neuromuscular disease [79]. Hence, in patients with low PI,max, the addition of sniff nasal
pressures further improved the diagnostic process and some patients were consequently
classified with normal respiratory muscle force [80]. The two techniques should hence be
considered complementary, rather than interchangeable. Normal values for the sniff
63
T. TROOSTERS ET AL.
nasal pressure are available [72]. Sniff measurements may be problematic in patients with
significant upper airway disease. Since the sniff is a very short manoeuvre, damping of
the pressure from the oesophagus to the mouth and nose may occur in patients with
obstructive lung disease, such as cystic fibrosis [76].
Much like the PI,max, the sniff nasal pressure reflects a global measure of inspiratory
muscle strength and not of diaphragm strength [74].
Equipment. Essentially the equipment can consist of the same pressure transducer as the
one used in the assessment of the PI,max. A perforated plug with a tube is used to occlude
the nostril. The tube is connected to the pressure transducer and the pressure–time curve is
recorded for inspection and quality control. The peak pressure is reported after a series of
maximal sniffs separated by normal breathing. A plateau is generally obtained after 5–10
sniffs. As the sniff pressure is a very brisk manoeuvre the recording of the trace should be
done with high resolution to allow detection of the peak pressure. Currently these devices,
and accompanying software, are commercially available.
Measurement in oesophagus or stomach
In rare clinical cases, and to answer specific research questions, it may be useful to
measure the pressure in the oesophagus or in the gastric area. In the oesophagus the
pressure (Poes) is a reflection of the pleural pressure (Ppl); the gastric pressure reflects the
abdominal pressure (Pabd). The difference between both pressures is the "transdiaphragmatic pressure" (Pdi), which is a more specific measure of diaphragmatic function.
To obtain these pressures a latex balloon catheter is put in place. Generally this is done
by swallowing a balloon catheter introduced in the nose, after application of a local
anaesthetic spray to the nasal mucosa and the pharynx. Double lumen catheters are
available for simultaneous measurements of pressure above and below the diaphragm
(Pdi). Balloons placed over the catheters are 5–10 cm long, have thin walls and are filled
with y0.5mL of air to allow proper transmission of the pressure into the catheter.
Catheter mounted microtransducers are an alternative to the "classical" balloon
catheters. These transducers are accurate, but measure pressure only at one spot. Hence
the measurement obtained may be a less precise reflection of the overall Poes. In addition,
these catheters are much more expensive [17].
These tests are perceived by many patients as rather uncomfortable, but the results give
probably the best estimate of the pressures generated by the respiratory muscles during
normal breathing, during exercise, or during static manoeuvres or sniffs. When the
balloon is positioned in the stomach, gastric pressure can also be recorded during cough.
Hence "cough" pressure is recorded (Pcough) [81]. In healthy subjects, Pcough was reported
to be superior to PE,max, and the lower limit of normal is set at 12.9 kPa (132 cmH2O) for
male and 9.5 kPa (97 cmH2O) for female subjects. Recently, Pcough were found to be a
useful addition in the diagnosis of expiratory muscle weakness. In a significant number of
patients with low PE,max, Pcough was reported normal. By contrast only a few patients
with normal PE,max exhibited low Pcough [81]. As a noninvasive variant of Pcough Chetta
et al. [82] recently introduced the "whistle" pressures, measured at the mouth. Subjects
were asked to perform a short, sharp blow as hard as possible from TLC through a
reversed paediatric inhaler whistle.
Nonvolitional tests of respiratory muscle function
Measurements of maximal voluntary inspiratory or expiratory pressures at the mouth,
nose, or even using balloon catheters to measure oesophagus or gastric pressures, are
64
RESPIRATORY MUSCLE ASSESSMENT
biased by the motivation of the patient to collaborate with the tests. Maximal effort is
sometimes difficult to ascertain because of lack of patient motivation, anxiety, pain or
discomfort, submaximal central activation, poor mental status or difficulties in
understanding the manoeuvres.
To overcome the issue of submaximal (voluntary) activation, investigation of the
diaphragmatic function can be done through electrical [83] or magnetic [84] stimulation
of the phrenic nerve. The diaphragm is exclusively innervated by the phrenic nerve (left
and right). This nerve passes superficially in the neck and can be stimulated relatively
easily. In addition, electromyography of the costal diaphragm can be carried out. When
the latter is done, the phrenic nerve latency can be studied [85, 86], which allows lesions of
the phrenic nerve to be detacted. Pressures developed after twitch stimulation of the
phrenic nerve can be measured transdiaphragmatically, or at the mouth. Although this
technique is not often used in clinical routine, there are specific situations in which it may
provide useful and unique information [87].
Respiratory muscle endurance
Although maximal in- and expiratory muscle strength gives important information on
respiratory muscle function, the respiratory muscles (especially the inspiratory muscles)
should be able to cope with endurance tasks. Measurements of respiratory muscle
endurance, therefore, give clinicians further insight in the function of the respiratory
pump, and may unmask early task failure. In the authors’ opinion, measurements of
inspiratory muscle endurance are especially helpful when inspiratory muscle weakness is
discrete, and its clinical consequence is unclear. In the clinic, respiratory muscle
endurance is generally assessed using one of the following techniques:
Maximal sustainable voluntary ventilation
The maximal sustainable voluntary ventilation (MSVV) is measured, or estimated
from protocols with incremental ventilation [88]. The achieved sustainable ventilation
is then reported as a fraction of the actually measured 12–15 s maximum voluntary
ventilation (MVV), and/or as a fraction of the predicted MVV. MSVV should be y60–
80% of the 12 second MVV. This test can be considered as a test of in- and expiratory
muscles, but it is relatively sensitive to changes in airway obstruction, and needs careful
control and adjustment of CO2 tension in arterial blood, by adding or removing dead
space or CO2 to the inspired air. In patients with severe airflow obstruction, MVV may
be low due to important dynamic compression of the airways during the vigorous 12 s
manoeuvre. Therefore, MSVV/MVV may seem relatively high in these patients, whereas
other measurements of endurance showed reduced respiratory muscle endurance in
COPD [89]. In a variant of this test proposed for COPD patients, patients are asked to
sustain a ventilation of 66–75% of their MVV [90]. This test allows comparison within
one subject, but normal values are not available.
Incremental threshold loading
Patients are asked to breath against increasing inspiratory loads. The inspiratory
threshold load is increased every 2 min [91]. The test can be compared with an
incremental exercise test. The highest pressure that patients can sustain for 2 min in the
incremental protocol is called maximum threshold pressure (Pthmax). Generally patients
should be able to reach a pressure equivalent to 75–80% of PI,max. Johnson et al. [92]
65
T. TROOSTERS ET AL.
reported that the Pthmax/PI,max was dependent on age. Important learning curves are
reported for this test, and the test should be repeated at least two to three times [93, 94]. One
study, conducted in COPD patients confirms the learning curve for the Pthmax at which
patients could continue breathing, but since PI,max showed a similar learning curves, the
Pthmax/PI,max ratio remained constant (61% in test 1 and 67% in test 4) [95]. Due to the
incremental nature of the test, however, it can be criticised as a straightforward measure of
endurance. Alternatively, the maximum sustainable threshold load can be determined. The
sustainable load is the load that can be sustained forw10 min. This technique reflects better
the concept of "endurance", but it is time consuming.
Recently, an expiratory incremental threshold loading test was developed, and used in
healthy subjects and subjects with COPD [32]. Interestingly, the authors reported that the
expiratory pressure that was achieved following an incremental protocol was only 44%
of PE,max in COPD. In healthy subjects 87% of PE,max was reached. The clinical
consequences of these findings may be illustrated by the recent finding that expiratory
muscle training in COPD may be a successful training strategy to improve exercise
capacity and dyspnoea in patients with COPD [33]. Further studies, however, should be
conducted to assess the usefulness of such an intervention on a larger scale.
Endurance time at a given threshold intensity
From the work of Nickerson and Keens [96], and others [91, 97] it can be deduced
that an inspiratory load of 60% of the PI,max can generally be sustained forw10 min. As a
simple test of respiratory muscle endurance, hence, patients can be asked to breath at a
fixed inspiratory load equal to 60% of PI,max. When subjects fail to continue breathing
against this resistance at any time point earlier than 10 min, respiratory muscle
endurance can be assumed impaired. Although easy to apply in clinical routine, this test
has many methodological problems that impair the use of this test in clinical studies. The
most important problem is probably the fact that the time to fatigue is related to
the breathing pattern (i.e. the inspiratory time (TI)/total respiratory time (Ttot) ratio).
The higher this ratio, the sooner fatigue will occur. Hence TI/Ttot should be carefully
controlled and maintained at y0.4 during the test [98]. Despite these methodological
shortcomings the present authors use this test as a useful addition to a measurement of
PI,max in patients presenting with muscle weakness. In this case the test may give
clinicians information on the susceptibility to inspiratory muscle fatigue. In patients with
normal inspiratory muscle strength, the test is considered of less clinical value, as the
pressures that should be sustained are far from those achieved in physiological
conditions.
Conclusions
The measurement of respiratory muscle force evolved from a technique used in clinical
physiology studies to a measurement that gained importance in the clinical routine.
Assessment of respiratory muscle force is extremely useful to understand the aetiology of
dyspnoea, and the detection of respiratory muscle weakness has consequences in the
treatment of patients. The most obvious example is the introduction of respiratory
muscle training in patients with respiratory muscle weakness. Measurement of
respiratory muscle strength is not restricted to patients with lung disease and should
also be carried out in neuromuscular, systemic and cardiologic disease. In addition, in the
follow-up of patients treated with drugs that may induce myopathy, the assessment of
respiratory muscle function is advised. In the large majority of cases the assessment of
66
RESPIRATORY MUSCLE ASSESSMENT
maximal inspiratory pressures give sufficient information to clinicians. In rare cases
measurements of pressures in the abdomen or oesophagus may be needed. In a limited
number of laboratories nonvolitional assessment of the respiratory muscles is done
through magnetical or electrical stimulation of the phrenic nerve.
Summary
Respiratory muscle weakness has serious clinical consequences. The assessment of
respiratory muscle function and the detection of respiratory muscle weakness has a
place in the clinical decision tree of many diseases, including lung disease,
neuromuscular diseases and others. Equipment to measure respiratory muscle
strength has become available and assessment of respiratory muscle force through
the assessment of maximal in- and expiratory pressures at the mouth (PI,max, PE,max),
has become a routine assessment in many lung function laboratories. In rare cases
more elaborate measurements, including transdiaphragmatic pressures, cough
pressures or measurements applying electrical or magnetical stimulation of the
phrenic nerve, can be helpful in the diagnostic process. Clinicians should be aware that
respiratory muscle force is approached indirectly by measuring the pressure generated
by the respiratory pump. The mechanics of the pump should be taken into account
when interpreting the results. Normal values are available, but large variability is
present. Part of this variability is explained by the methodological differences
described in this chapter.
Nevertheless, since respiratory muscle weakness can be treated in many cases by
respiratory muscle training, or tapering of treatment with drugs that may induce
respiratory muscle weakness (e.g. corticosteroids) or may help clinicians decide on
mechanical ventilation strategies, knowledge of respiratory muscle dysfunction opens
a window of clinical treatment opportunities. Hence, properly performed assessment
of respiratory muscle function should be possible in any well-equipped lung function
laboratory.
Keywords: Dyspnoea, muscle force, respiratory muscle.
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71
CHAPTER 5
Forced oscillation technique and impulse
oscillometry
H.J. Smith*, P. Reinhold#, M.D. Goldman}
*Research in Respiratory Diagnostics, Berlin, Germany. #Friedrich-Loeffler-Institute, Jena, Germany.
}
David Geffen School of Medicine, University of California, Los Angeles, USA.
Correspondence: H.J. Smith, Research in Respiratory Diagnostics, Bahrendorfer Str. 3, 12555 Berlin,
Germany.
Conventional methods of lung function testing provide measurements obtained during
specific respiratory actions of the subject. In contrast, the forced oscillation technique
(FOT) determines breathing mechanics by superimposing small external pressure signals
on the spontaneous breathing of the subject [1]. FOT is indicated as a diagnostic method
to obtain reliable differentiated tidal breathing analysis. Because FOT is performed
without closure of a valve connected to the mouthpiece, and without maximal or forced
respiratory manoeuvres, it is unlikely that FOT itself will alter airways smooth muscle
tone [2].
FOT utilises the external applied pressure signals and their resultant flows to determine
lung mechanical parameters. These pressure–flow relationships are largely distinct from
the natural pattern of individual respiratory flows, so that measured FOT results are, for
the most part, independent of the underlying respiratory pattern. Therefore, oscillometry
minimises demands on the patient and requires only passive cooperation of the subject:
maintenance of an airtight seal of the lips around a mouthpiece and breathing normally
through the measuring system with a nose-clip occluding the nares. Potential
applications of oscillometry include paediatric, adult and geriatric populations,
comprising diagnostic clinical testing, monitoring of therapeutic regimens, and
epidemiological evaluations, independent of severity of lung disease. Oscillometry is
also applicable to veterinary medicine.
The last two main sections Relevance of oscillometry in clinical practice and
Oscillometry in the clinical pulmonary laboratory emphasise clinical aspects of
application and interpretation of FOT rather than methodological details and
technological solutions, which are discussed in the two sections that follow immediately
below. Clinical application of FOT does not require mastery of the mathematical
infrastructure of the technical methodology, and readers interested in the clinical use of
FOT may find it more useful to begin with these clinical sections and refer subsequently
to methodological and technological details.
Methodology of impulse oscillation technique
The mechanical basis of oscillometry concerns use of external forcing signals, which
may be mono- or multifrequency, and applied either continuously or in a time-discrete
manner. The impulse oscillation technique is characterised by use of an impulse-shaped,
time-discrete external forcing signal.
Eur Respir Mon, 2005, 31, 72–105. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
72
OSCILLOMETRY: FOT AND IOS
The most useful aspect of applying FOT pressure pulses rather than pseudo-random
noise (PRN) is improved time resolution of the measurement. The impulse oscillation
technique allows measurement of up to 10 impedance spectra per second. This permits a
useful analysis of intra-breath variation in impedance, comparable to that obtained with
mono-frequency applications. However, a disadvantage of such high impulse rates is the
inability to record longer respiratory time constants that may be more informative
concerning respiratory abnormalities. For this reason, the common application of
impulse oscillation utilises recordings of 5 impedance spectra (5 impulses) per second.
An additional benefit of impulse oscillation is the simplicity of the hardware needed to
generate the forced signals, allowing smaller, more efficient electronic and mechanical
structures with minimal power loss.
A unique aspect of applying pulses of pressure to the respiratory system is the fact that
the entire energy of all applied pressure harmonics is applied within a very short time
period. This causes a higher impact to the respiratory system compared with sinusoidal
or PRN applied pressures, and may be perceived by some patients as a slightly
unpleasant respiratory sensation during measurement.
Peculiarities of aperiodic waveforms
The impulse oscillation method applies aperiodic waveforms using an impulse
generator that produces pressure pulses of limited magnitude and 30–40 ms duration.
These pulses define specific amplitudes and phases of the inherent sinusoidal
components. The time-course of such practical pressure pulses applied to the respiratory
system is not a true Dirac-impulse, defined as having virtually infinite magnitude and
infinitesimal time duration, which would provide a continuous spectrum of frequencies
with the same amplitude. Thus, the terms "impulse-shaped" oscillations and "impulse
pressures" are used to indicate realistic practical pressure pulses rather than
mathematically defined impulses.
The short duration of the impulse-shaped waveform itself provides linearity of
pressure and flow signals in the face of within-breath dynamic changes in the respiratory
system. In contrast, the longer time needed with PRN to embed a range of periodic
functions decreases time resolution, resulting in increased noise of calculated impedance
related to any time-dependence of dynamic changes in the respiratory system.
The characteristic feature of any aperiodic waveform is the resulting continuous
spectrum after transformation of its time course into the frequency domain, using a
Fourier integral rather than a Fourier series, in the fast Fourier transform (FFT).
Thus, the advantage of continuous spectra is particularly important in abnormal
respiratory systems with regional nonhomogeneities (fig. 1), where resistance, reactance
and coherence spectra may manifest deviations from the normally smooth and uniformly
continuous spectral courses.
In contrast, spectra resulting from FFT analyses of periodic multifrequency forcing
such as PRN [3] are discontinuous. Discrete values of impedance are obtained with a
frequency resolution determined by the frequencies of the included sinusoidal
components. As a result, the course of such discrete spectra often may require postprocessing to smooth the PRN spectra [3]. To improve interpretation of discrete PRN
spectra, it is common to approximate the overall spectral range by smoothing with linear,
quadratic or logarithmic functions. However, such smoothing inherently diminishes
information contained in characteristic peaks and plateaus of impedance that may
otherwise provide insight into superimposed parallel resonance phenomena, e.g. related
to upper airway constriction.
Impulse power spectra of pressure and flow generated by the impulse oscillometry
73
kPa·s·L-1
H.J. SMITH ET AL.
Rrs
0.5
Xrs
0.3
0.4
0.2
0.3
0.1
0.2
0
Coherence g2 (ƒ)
Regional
nonhomogeneity
ƒres Reactance Xrs(ƒ)
0.9
0.8
Resistance Rrs (ƒ)
0.7
Reactance area (AX)
0.1 -0.1
0 -0.2
g2
1.0
0
5
10
15
20
Frequency ƒ Hz
0.6
25
30
35
0.5
Fig. 1. – Representative data for spectra of respiratory resistance (Rrs(f); ––), reactance (Xrs(f); - - -) and
coherence (c2(f); - – - – ) are plotted, between 3 and 35 Hz, for a normal adult, during forced oscillation using
pulse-shaped forcing generated by the impulse oscillometry system. Resonant frequency (fres) is shown where the
Xrs(f) tracing crosses zero. The shaded area, below zero Xrs and above Xrs(f) tracing between 5 Hz and fres, is
the integral of Xrs(f) from 5 Hz to fres, and is designated the reactance area (AX). Regional nonhomogeneities
may manifest deviations from the normally smooth and uniformly continuous spectral courses.
system (IOS) are shown in figure 2 over a frequency range of 0.1–35 Hz. The energy
distribution provides practical assessment of low (ƒ5 Hz), as well as high (w20 Hz)
frequency ranges, with decreased amplitude at higher frequencies to minimise nonlinearities due to acceleration of the moving air column [4]. Enhanced amplitudes at the
5
Attenuation dB
0
-5
-10
-15
-20
-25
0
5
10
15
20
Frequency ƒ Hz
25
30
35
Fig. 2. – Power spectra of flow (––), and pressure (----), for the discrete pulse-shaped forcing signal generated by
the impulse oscillometry system. Spectra plotted at frequencies 0.1–35 Hz. Pressure and flow power are highest
at 3–20 Hz. Less power is needed at frequencies w20 Hz, because "competing" higher harmonics of the patient’s
respiratory flow are very small at these frequencies.
74
OSCILLOMETRY: FOT AND IOS
lower frequencies limit the influence of higher harmonics of spontaneous breathing
frequencies.
Impulse oscillometry system
The IOS measuring-head (fig. 3) is functionally similar to PRN systems designed for
the determination of input impedance [1, 5–12]. The characteristic feature of the IOS [13]
is the generation of recurrent aperiodic impulse-shaped forcing signals of alternating
direction.
Flow is measured by a Lilly-type heated screen pneumotachograph with a resistance of
36 Pa?s?L-1, providing a common-mode rejection ratio ofw60 dB up to 50 Hz [12, 14, 15]
for the combination of pneumotachograph and flow transducer system. At flow rates
v15 L?s-1 the heated pneumotachograph is linear within 2%. The proximal side of the
pneumotachograph is connected to a pressure transducer. To guarantee suppression of
technical influences and to avoid phase differences, both pressure and flow channels use
matched transducers of the same type, SensorTechnics SLP 004D [16]. Pressure and flow
signals are sampled at a frequency of 200 Hz and digitally converted by a 12-bit
analogue-to-digital converter. Analogue anti-aliasing low-pass filtering is realised by a
fourth-order Bessel filter at a border frequency of 50 Hz, providing a damping at Nyquist
frequency of y75 dB.
Tuning of the IOS impulse generator involves both volume displacement of the
loudspeaker membrane and magnitude of the terminating resistor. The terminal resistor
RS 232 to computer
DSP
Loudspeaker
Impulse generator
Y-adapter
Metal screen
w
flo
lse flow
pu
y
Im ator
ir
sp
Re
Terminal
resistor
Pneumotachograph
(heated)
Flow
transducer
Mouthpiece
Pressure transducer
Fig. 3. – Schematic diagram of the impulse oscillometry system measuring-head and connectors: loudspeaker
enclosure at top, connected to a y-adaptor at one upper arm, an exit for flow with terminal resistor at the
second upper arm and a lower arm connected to the pneumotachograph. A mouthpiece is connected to the
open side of the pneumotachograph. Pulsatile flows generated by the loudspeaker are shown as a lightly shaded
thick line, part of which exits through the terminal resistor and part of which flows through the
pneumotachograph and mouthpiece. The patient’s normal respiratory flow is shown as a shaded thick line from
the mouthpiece though the Y-connector, exiting through the terminal resistor. The resistance of the terminal
resistor is 0.1 kPa?s?L-1 and the deadspace of the Y-connector/pneumotachograph/mouthpiece is y70 mL.
75
H.J. SMITH ET AL.
provides a low-impedance pathway for normal respiratory flow, which, at the same time,
is high enough to minimise loss of energy of superimposed impulses so that sufficient
impulse pressure is transmitted into the respiratory tract. Both components determine
the linear working range of the unit and range of input impedances, which can be
measured to maintain international recommendations [16, 17]. A terminal resistor of
y0.1 Pa?s?L-1, in combination with a volume displacement of 40 mL, which is
accelerated by the loudspeaker membrane in v40 ms, results in maximal peak-to-peak
impulse pressures of 0.3 kPa and minimises interference of underlying higher harmonics
of respiratory frequency that contribute "noise" to the oscillatory pressure and flow
signal [10].
International recommendations for electromechanical performance are maintained in
the IOS by use of advanced transducer technology and global mean spectral data derived
from IOS are generally comparable to those obtained by the pseudo-random noise
method of Làndser [6], Delecourt et al. [8] and Skloot et al. [18].
The measurement is performed as follows: while the subject spontaneously breathes
ambient air via the tubing between mouthpiece and terminating resistor, the loudspeaker
generator transmits brief pressure impulses via Y-adapter, pneumotachograph and
mouthpiece into the respiratory tract; pneumotachograph and pressure transducer
register the composite signals containing breathing activities and the forcing impulse
signal for further processing.
Further processing of digitised impulse data. Flow and pressure channels contain both
the underlying respiratory system flow and pressure, and the superimposed forced
oscillation signals.
By definition, respiratory input impedance is the transfer function or ratio of effective
pressure (Prs) and flow (V9rs), derived from the superimposed forced oscillations, after
being discriminated from underlying respiratory pressure and flow and their harmonics.
In the mathematical sense, all components are considered "complex", characterised by
modulus and phase.
Prs (f )
ð1Þ
Multifrequency input impedance Zrs(f )~ |0
f0 < f ƒf maxg
V rs (f )
|
Discrimination of superimposed forced oscillations from the underlying respiratory
pressure and flow in the IOS is focused on individual impulses, based on pressure and
flow sampling intervals that include both the impulse stimulus and the respiratory
system reaction to the impulse. Figure 4a gives an example of such a sampling
interval for flow.
A "baseline" straight line segment is inserted between the start- and end-points of
separate sample segments of both pressure and flow. The baseline is a simple linear
approximation of the underlying respiratory flow and pressure throughout the single
impulse excitation. Baseline approximation has proved to be a useful and reliable
technique to decrease the respiratory component of the composite signals for pressure
and flow. Spline reconstruction, sinusoidal approximation of underlying respiratory
pressure and flow or high-pass digital filtering were not as useful. Baseline correction and
offset elimination, as can be seen in figure 4b, allow rectangular windowing prior to the
FFT to effectively reduce spectral leakage and improve the signal-to-noise ratio.
Resolution of calculated pressure and flow spectra is increased by adding numerical
zero values to the real sampling points of the corrected primary data segment. This
procedure allows formation of exactly 2n samples, compatible with FFT requirements.
Choice of sampling interval as well as addition of zero values to this interval allow
76
OSCILLOMETRY: FOT AND IOS
Segment 160 ms
0.75
0.50
Approximated
baseline
0.25
0.00
Pulserate 5·s-1
200 ms
-0.25
-0.50
Corrected impulse flow L·s-1
b)
0.0
0.1
0.4
0.2
0.3
0.4
Recording time s
150
Resulting impulse pressure
0.2
0.7
0.6
0.5
75
Reactive response of respiratory system
0.0
-0.2
Duration
of generator
impulse
0
30
0
Baseline
Impulse flow
60
90
Sampling time ms
120
150
Corrected impulse pressure Pa
Amplitude of primary flow L·s-1
a) 1.00
-75
Fig. 4. – a) Primary flow recording. Note that the primary flow includes the patient’s expiratory respiratory flow
with pulses of "loudspeaker flow" superimposed. The dash-dot line shows the linearly approximated respiratory
flow if there were no superimposed pulsatile flow from the loudspeaker included. By referring all pulsatile flow
and with a similar process utilised for the continuous pressure tracing to the approximated baseline, the flow
and pressure related purely to the pulse generated by the loudspeaker, with the patient’s respiratory flow and
pressure subtracted from the total pressure and flow signals, may be derived. b) Corrected impulse tracings of
flow (––) and pressure (----), derived from baseline correction shown in figure 4a, as prepared for input into the
fast Fourier transform. Duration of impulse corresponds to actual movement time of the loudspeaker diaphragm
(y40 ms in duration). Initial flow response of the respiratory system begins almost immediately. After
loudspeaker movement has ended, the respiratory system continues to respond in a "reactive" manner, with
pressure reaching its peak and then declining, while flow falls below baseline, and then gradually returns to
baseline.
adjustment of spectra concerning low-end border frequency as well as numerical
resolution.
To further improve quality of calculated impedance data, impulses that do not fulfill
defined reliability criteria are rejected. The critical segments of respiration are the phase
transitions between inspiration and expiration. At these zero-crossings of pressure and
flow, gradients of pressure and flow versus time are maximal, and the following principles
are implemented to establish reliability. Slopes of baseline corrections for pressure or flow
w0.7 indicate dominance of the underlying respiratory pressure–flow pattern, and the
77
H.J. SMITH ET AL.
impulse is rejected because separation of the superimposed impulse pressure and flow is
not reliable. Absolute peak flow within the impulse segment must exceed 0.02 L?s-1.
Pulses with less flow are rejected because of very small flow values in the denominator of
the input impedance Equation (1) [19] and the resulting mathematical errors. Finally,
impulses that yield negative resistance values at any frequency after FFT are rejected.
Impulse rate and sampling interval. Impulse rate and selected sampling interval effects
on calculated impedance have been evaluated in vitro as well as in calves of comparable
size to adult humans [20, 21]. No significant effect of impulse rate was observed between 1
and 5 impulses per second.
In contrast, different sampling durations led to significantly different low-frequency
respiratory system reactance (Xrs) results. Using 16 sampling points (80 ms) for data
analysis, no useful information was obtained for Xrs ƒ5 Hz. Use of 32 sampling points
(160 ms) per impulse for data analysis provided useful Xrs5 data. However, sampling
durations of 320 ms did not improve impedance results and coherence c2 ƒ5 Hz was
significantly decreased, consistent with deterioration of calculated impedance quality due
to interactions with spontaneous breathing signals. These findings underlie the current
recommendation to use sampling intervals of 32 sampling points per impulse,
corresponding to an optimised impulse rate of either 3 or 5 impulses per second.
Coherence. The coherence function is defined as the square of the cross-power spectrum
divided by the product of the auto-spectra of pressure and flow at any forcing frequency.
Ranging between 0 and 1, it is a measure of the available linearity [22].
jGV 0 P (f )j2
f0 < f ƒf maxg
ð2Þ
GV 0 V 0 (f ):GPP (f )
2
É et al. [6] found that when the coherence c i0.95, PRN impedance data
Làndser
show a coefficient of variation CV%v10%. Subsequently, coherence value thresholds
of 0.9 or 0.95 have been widely used to accept FOT data. However, if methods other
than PRN forcing input signals or data acquisition are used, this original threshold
value of coherence is not applicable. Miller and Pimmel [23] showed that estimated
variance of calculated impedance is a function of coherence and the number of
estimates averaged. Use of pseudo-random noise techniques commonly includes three
measures of impedance of 16 s each [5]. Three replicate measures require coherence c2
w0.9 to yield an estimated standard error of 10%. In contrast, IOS commonly
includes the average of w100 separate FFT analyses and, accordingly, requires less
perfect coherence for the average data to provide an estimated standard error of
v10%. For clinical purposes it is recommended to use a coherence value of c2 i0.6 at
5 Hz for the acceptance threshold, provided that at least 100 FFT analyses are
averaged. Coherence improves as a function of oscillatory frequency to c2 i0.9 at
10 Hz and higher frequencies (fig. 1).
c2 (f )~
Interpretation of oscillation mechanics
Monofrequency oscillations provide a measure of total respiratory impedance (Zrs)
that includes airway resistance, and elastic and inertive behaviour of lungs and chest wall
at one oscillation frequency. In contrast, multifrequency oscillation methods, such as
pseudo-random noise or impulse oscillation provide measures of respiratory mechanical
properties in terms of Zrs, as a function of frequency (f), allowing the recognition of
characteristic respiratory responses at different oscillation frequencies.
78
OSCILLOMETRY: FOT AND IOS
Zrs has been described in engineering terms by so-called "real" and "imaginary"
components. In medical use, preferred terms are resistance (Rrs) and reactance (Xrs),
respectively.
Zrs(f )~Rrs(f )zjXrs(f ) f0 < f ƒf maxg
ð3Þ
|
In contrast to Rrs, concepts of Xrs are not yet widely appreciated, largely because of
greater complexity of reactive parameters, as well as their numerical characteristics.
The present discussion emphasises the importance of consideration of both Xrs as
well as Rrs for interpretation of respiratory mechanical properties.
Most commonly, the oscillation frequency scale utilised for multifrequency oscillation
methods includes about 5–30 Hz [7, 17, 24]. Oscillation frequenciesv5 Hz are affected by
higher harmonics of underlying natural respiratory frequency [25]. Higher oscillation
frequencies are increasingly affected by "shunt" properties of the upper airways [6, 13, 26]
and capacitive impact of face masks [27] when used.
Respiratory resistance
The resistive component of respiratory impedance, Rrs, includes proximal and distal
airways (central and peripheral), lung tissue and chest wall resistance. Normally, central
resistance dominates, depending on airway calibre and the surface of the airway walls,
while lung tissue and chest wall resistance are usually negligible [7, 28].
Rrs may be considered within normal limits if Rrs at 5 Hz (Rrs5) is within ¡1.64 sd of
the predicted value. Rrs5 values between 1.64 and 2 sd above predicted may be considered
minor, w2 sd moderate and w4 sd above predicted severe obstruction.
In previous reports the calculation of per cent predicted has been used. Rrs5 values that
did not exceed 150% predicted, defined in bronchial challenge comparing changes in Rrs5
to a 20% decrease in forced expiratory volume in one second (FEV1) and 50% increase in
airway resistance (Raw), were considered within normal limits [29–31]. However, it is now
recognised that a 20% decrease in FEV1 is a substantial abnormality, and normal limits
for Rrs might be more profitably defined in their own right, without requiring a specific
relationship to arbitrary spirometric criteria.
In healthy subjects, Rrs is almost independent of oscillation frequency [7, 32, 33], but
may increase slightly at higher frequencies due to the upper airways shunt effect [6, 26].
When proximal (central) or distal (peripheral) airway obstruction occurs, Rrs5 is
increased above normal values. The site of airway obstruction is inferred from the
pattern of Rrs, as a function of oscillation frequency, adjusting as necessary for subject
age [2, 25, 34–40]. Proximal airways obstruction elevates Rrs evenly independent of
oscillation frequency [32].
In distal airways obstruction, Rrs is highest at low oscillation frequencies and falls with
increasing frequency. This negative frequency-dependence of Rrs has been explained in
terms of intrapulmonary gas flow redistribution, due either to peripheral pulmonary
nonhomogeneities or to changes in peripheral elastic reactive properties [6, 8, 34, 35]. As
peripheral resistance increases, Rrs becomes more frequency dependent [26, 36–38].
Frequency dependence of Rrs may be a normal finding in small children [39, 40] and may
be greater than in adults in the presence of peripheral airflow obstruction [8].
Respiratory reactance
The reactive component of respiratory impedance, Xrs incorporates the mass-inertive
forces of the moving air column in the conducting airways, expressed in the term
inertance (I) and the elastic properties of lung periphery, expressed in the term
79
H.J. SMITH ET AL.
capacitance (Ca).
1
v~2:p:f f0 < f ƒf maxg
ð4Þ
Xrs(f )~v:I{ :
v Ca
Most importantly, in medicine, it is emphasised that respiratory Ca is not identical to
compliance. The component of Xrs associated with Ca is defined to be negative in
sign. It is most prominent at low frequencies. In contrast, the component of Xrs
associated with inertance is always positive in sign and dominates at higher
frequencies. Thus, interpretation of Xrs is primarily influenced by the oscillation
frequency range under consideration.
Low frequency capacitive Xrs essentially expresses the ability of the respiratory tract to
store capacitive energy, primarily resident in the lung periphery. In both fibrosis and
emphysema this ability is reduced: in fibrosis because of the stiffness of the lung; in
emphysema because of hyperinflation and loss of lung elastic recoil. Distal capacitive
reactance at 5 Hz (Xrs5) manifests increasingly negative values either in restriction or in
hyperinflation. Thus, Xrs5 characterises the lung periphery, but is nonspecific as to the
type of limitation. Additional information is needed to differentiate peripheral
obstruction from peripheral restriction. This is not usually problematic in clinical
practice.
Capacitive and inertive elements have been modelled by a number of authors [2]. Many
simplifications have been attempted that include serial and parallel circuit elements, to
define an approximation of a normal Xrs spectrum with a zero-crossing exactly at
resonant frequency and positive slope throughout. The frequency range utilised in
multifrequency oscillometric forcing signals should allow for determination of the serial
resonance of the respiratory system under investigation [5].
Resonant frequency. Resonant frequency (fres) is defined as the point at which the
magnitudes of capacitive and inertive reactance are equal.
1
v0 :I~ :
v0 ~2:p:f res
v0 Ca
ð5Þ
1
pffiffiffiffiffiffiffiffiffiffi
ð6Þ
:
:
2 p I :Ca
Because fres can vary over a considerable range and, thereby, appear in close
proximity to oscillation frequencies dominated by either capacitative or inertive
properties, this parameter should not be directly interpreted in terms of a particular
mechanical property of the respiratory system. However, it is a convenient marker to
separate low-frequency from high-frequency Xrs. Thus, low-frequency capacitive Xrs
is dominant at oscillation frequencies below fres, while high-frequency inertive Xrs is
dominant at frequencies above fres. In normal adults, fres is usually 7–12 Hz [33]. In
healthy young children, fres is larger than in adults and increases with decreasing age.
In respiratory disease, both obstructive and restrictive impairments of the distal
respiratory tract cause fres to be increased above normal [7, 26, 36]. The relevance of fres is
normally revealed in within-individual trends over time during bronchial or therapeutic
challenge.
f res~
Capacitive spectral range of reactance. Thoraco-pulmonary elasticity is commonly
viewed as a static property, and normally is investigated in the absence of resistive or
inertive mechanical losses. However, in oscillometry, capacitive Xrs is believed to
comprise useful information concerning peripheral airways mechanical properties. In
80
OSCILLOMETRY: FOT AND IOS
practice, determination of peripheral capacitive Xrs is determined at the lowest frequency
that is not highly interfered with by fundamental respiratory frequency and its harmonics
[41]. In the IOS method, Xrs5 is commonly utilised. In children breathing at high
respiratory frequencies, reactance at 10 Hz may be useful [42] in following bronchial and/
or therapeutic challenge.
Interpretation of Xrs5, is clearly different from that of conventional lung function test
parameters and, in particular, lung compliance. One striking feature of Xrs5 is its negative
value. The minus sign is derived from a general definition in natural sciences to
differentiate elastic properties from moments of inertia, the latter of which is always
positive. Therefore, the minus sign simply confirms a relationship with elastic properties.
Definition of abnormality has previously been based on increased negative readings
related to expected normal values of Xrs5. A differencew0.15 kPa?s?L-1 is most definitely
agreed to imply abnormal lung function, independent of Rrs, although abnormal lung
function may be present with smaller differences.
Capacitance versus dynamic pulmonary compliance. Dynamic pulmonary compliance
is derived from the relationship between oesophageal pressure and changes in lung volume
[43]. In contrast, oscillometry assesses the elastic properties of the respiratory system from
the out-of-phase relationship of simultaneously recorded transrespiratory pressure (Prs)
and central flow signals (V9rs), measured from superimposed oscillations and transformed
into Xrs. Therefore, the term capacitance, Ca, an equivalent of capacitive phase
information between the primary signals Prs and V9rs should be used. The frequency range
for such measures is always below fres.
In pulmonary fibrosis, dynamic lung compliance (CLdyn) is decreased below normal. In
a similar manner, oscillometry yields a decreased estimate of Ca, due to negative
displacement of low frequency Xrs. Both dynamic lung compliance and Ca reflect elastic
limitation and they trend together. Previous oscillometry studies have reported less
sensitivity than dynamic lung compliance in the early stages of restrictive disease [7, 17].
In contrast, pulmonary hyperinflation is associated with loss of lung elastic recoil and
increased CLdyn. Because of the loss of lung elastic recoil, peripheral airways are not
supported externally by lung recoil and the resultant partial peripheral airway
obstruction prevents the applied oscillometric signals from reaching peripheral
compliant areas. In this way, loss of lung elastic recoil indirectly causes a decrease in
Ca, with associated increased negative magnitude of low frequency Xrs, [26, 36]. Indeed,
low frequency Xrs is particularly sensitive to pulmonary hyperinflation, and while Rrs5
may be nearly normal or only moderately increased, Xrs5 is highly abnormal [44].
Reactance area. The index designated reactance area (AX) is a quantitative index of total
respiratory reactance Xrs at all frequencies between 5 Hz and fres. The integration of these
negative values of Xrs [45] creates an area between the reactance zero axis and Xrs,
providing an integrative function to include changes in the magnitude of low-frequency
reactance Xrs, changes in fres and changes in the curvature of the Xrs(f)-tracing. It is
represented graphically in figure 1 as the area under the zero axis of the reactance graph
above the Xrs(f) tracing.
fð
res
AX ~
Xrs(f ):df
ð6Þ
5
This integrative index provides a single quantity that reflects changes in the degree of
peripheral airway obstruction and correlates closely with frequency dependence of
resistance [18].
81
H.J. SMITH ET AL.
Inertive spectral range of reactance. During resting breathing at normal respiratory
frequencies, transrespiratory pressure is dissipated in resistive and elastic losses, whereas
inertive pressure losses are negligible [46]. In contrast, during forced oscillation, when
oscillatory frequencies are more than 10-times greater than normal respiratory frequency,
inertance (I) contributes significantly to dissipation of the externally applied pressures [16,
46]. As noted above, the inertive spectral range of Xrs is at frequencies above fres. These
frequencies reflect mechanical properties of the proximal conducting airways. However,
specific clinical interpretation of Xrs in this range is limited due to the wide variety of
influences that may appear in the upper respiratory tract and resultant resonant effects
that may change Xrs unpredictably. Changes in central airway calibre correlate more
strongly with resistive parameters and are less well represented in the inertive spectral
range.
Finally, it should be noted that oscillometric reactance occasionally demonstrates a
low-to-mid frequency plateau in the Xrs (f) tracing, which is suggestive of possible upper
airway obstruction [47–50].
Variability of oscillometric parameters
The majority of oscillometric parameters can usefully be assessed using the coefficient
of variation (CV%). Short-term variability should not exceed 10%, for magnitude of Zrs
and Rrs at frequencies i5 Hz [17]. Variability of Xrs is larger, because of physiological
and numerical characteristics. Xrs can be positive or negative and is commonly close to
zero. As a result, the calculation of CV% is not suitable to estimate the variability of most
of the Xrs values. Therefore, it is recommended to estimate variability of Rrs and Xrs
using standard deviation, 95% percentiles for normally distributed values or calculating
the absolute difference (range) between minimum and maximum of the Xrs parameters
[20, 51].
Methodological versus biological variability. Low methodological variability in
measures of impedance (Zrs) has been shown using physical models with rather high
reproducibility [20].
Biological variability is much more complex and incorporates intra-breath and intraand inter-subject variability. Intra-subject variability of consecutive measurements
within a specified time period, including circadian variability, day-to-day variability and
variability associated with changes in diseased airways has been reported previously [52–
56].
Intra-breath variability is of specific physiological interest. Commonly, Zrs is
determined as an average over a number of consecutive breathing cycles within 15–
30 s. However, flow- and volume-dependence of Rrs and Xrs within each breathing cycle
may be apparent during both inspiration and expiration, and have been shown in a
number of investigations to reflect specific pathophysiological characteristics [57–61].
Special application of impulse oscillometry system to animals
Application of the impulse oscillation technique has been described in different animal
species. The standard IOS device, originally developed for humans, has been validated
carefully in calves aged up to 6 months and weighing 35–150 kg [20, 27, 62] and in pigs
[63, 64]. For large animals, such as horses or adult cattle, a specially designed IOS unit
has been developed, which is capable of analysing larger flow and volume characteristics.
While no data are available in adult bovines, methodological validation and clinical
application of IOS in horses has been reported [65–67] and is still in progress.
82
OSCILLOMETRY: FOT AND IOS
Oscillometry in spontaneously breathing conscious animals requires the imposition of
applied pressure signals via a rigid mask. Without correcting measured impedance for the
facemask, the limit of frequencies that can be clinically analysed is v15 Hz. Higher
frequencies are substantially influenced by the capacitance of the facemask itself [27].
The most useful frequency range for clinical evaluation of respiratory impedance
differs in different species, primarily dependent on animal size. The lower the specific
frequency is, the more sensitive the measurements to the periphery of the respiratory
system are. For example, while the resonant frequency is between 5–12 Hz in calves,
depending on body weight, it isv5 Hz in horses. Accordingly, frequencies that reflect the
lung periphery in horses are lower than in calves.
In agreement with results in human medicine, peripheral airway obstruction is
characterised by a marked increase in magnitude of low frequency respiratory reactance
(|Xrs|) and in fres. In addition, the resistance spectrum of Rrs shows increased negative
frequency dependence and increase in low frequency Rrs (v5 Hz). Upper airway
narrowing is characterised by a parallel increase in Rrs at all frequencies with no change
in the frequency dependence of Rrs. No significant changes occur in reactance with upper
(large) airway narrowing.
Relevance of oscillometry in clinical practice
This section offers a perspective on clinical use of the FOT method of determining Zrs
that may be summarised as follows:
1) FOT provides useful clinical information that prominently includes functional
assessment of small, peripheral airway behaviour beyond that available from
commonly used pulmonary function tests (PFT).
2) Because of its sensitivity to peripheral airway function, FOT in its own right, apart
from other PFT results, provides useful guidance in clinical patient management.
3) The prominence of peripheral airway functional assessment provided by FOT
derives both from Xrs as well as Rrs.
4) The importance of Xrs is amplified by recognition of different Xrs-characteristics at
low, i.e. below resonant frequency (vfres), and high (wfres) oscillation frequencies.
The last issue is considered in detail in the foregoing technological sections.
Briefly, it is noted here that original technical descriptions of FOT included calculation
of the magnitude (|Zrs|) and phase (Q) of Zrs, i.e. polar coordinates [1, 5]. This engineering
description gave way to clinical research descriptions of Rrs and Xrs in Cartesian
coordinates. As noted in the technology section, Xrs relates to peripheral airway
properties at oscillation frequencies vfres, and to central conducting airways at
frequencies wfres. The use of the magnitude of Xrs (|Xrs|) rather than Xrs itself in
algebraic manipulation of reactance data is emphasised, because this increases sensitivity
to changes in peripheral airway mechanical properties, as demonstrated later in this
section in reference to previously reported studies.
The section that follows includes discussion of monofrequency, pseudo-random noise
and pulse-shaped pressure oscillations, as these three methods are currently in common
use. The general principles apply to all methods of FOT for the most part where issues
concern specific use of multifrequency rather than monofrequency or IOS rather than
PRN this difference is stated explicitly. Previous published reviews have discussed
theoretical and modeling aspects of FOT [2, 45, 68]. The present discussion does not
include an exhaustive review of clinical research investigations from the Barcelona group
[69–73], Leuven [6, 29, 32, 36, 39, 74–78], London [7, 79–81], Paris [8, 82–86] and
83
H.J. SMITH ET AL.
Vandoeuvre-les-Nancy [25, 87–91], which have all helped to provide the essential
infrastructure for clinical application of oscillometry. Instead, the authors focus on
published studies in relation to current work that permit practical establishment of
oscillometry in the routine clinical pulmonary function laboratory. Furthermore,
because the authors have worked more intimately with IOS, illustrative examples of
current work with this technology are included.
The relative advantages of each type of FOT may be summarised by noting that the
simplest form for clinical practitioners is monofrequency sinusoidal pressure application
[2, 68]. Measurement of Rrs with this method is applicable to patients with sleep apnoea
and those using continuous positive airway pressure or mechanical ventilation.
Multifrequency FOT provides further characterisation of respiratory mechanics,
including variation of Rrs and Xrs with oscillation frequency. PRN FOT has been applied
to the description of patients with asthma, bronchitis, emphysema, diffuse interstitial
lung disease and thoracic wall deformities, and to assess bronchial or therapeutic
challenge [6, 29, 32, 36, 74–78]. Multifrequency FOT using PRN imposes a more gentle
forcing signal perturbation than IOS, and has not been noted to provoke
bronchoconstriction. IOS differs from PRN by utilising brief pressure pulses of 30–
40 ms duration. These pulses result in respiratory pressure responses that may be
perceived as a slightly unpleasant respiratory sensation in some subjects. The brief
pressure pulses provide convenient time-trend analyses and within-breath changes of Rrs
and Xrs not available with PRN. IOS is most familiar to the current authors, and
therefore, this and the following section relate current clinical work with IOS to
previously published reports of both PRN and IOS.
The clinical relevance of FOT may be assessed, as with any test of physiological
function, in terms of its utility in diagnosis. Two general approaches are considered: first,
the use of FOT as part of an initial complete diagnostic evaluation, including spirometry,
body plethysmography, and gas distribution and exchange measures; secondly, the use of
FOT as a means of monitoring response to treatment can include both bronchial and
therapeutic challenge.
Summary information is presented here concerning the utility of FOT in assessing
severity of lung disease, degree of airway reactivity, reversibility of airflow obstruction,
and stratification of breathing mechanics between central and peripheral airways.
Current clinical relevance of FOT relates significantly to the broad range of patients
that may conveniently be evaluated. In contrast to standard PFT requiring maximal
coordinated efforts, FOT requires only normal quiet breathing with the lips tightly closed
to avoid airflow leak, and the wearing of a nose-clip. For this reason, children can be
easily studied, often as early as 3 yrs [8, 19, 92–94]. Similarly, elderly subjects, those with
severe airflow obstruction or those with neuromuscular disease who find maximal forced
respiratory efforts difficult to perform are able to breathe normally for FOT testing [95–
98]. Portability of commercial FOT instruments permits lung function testing at the
bedside or, for occupational lung disease studies, at the place of work [18].
FOT places minimal performance demands upon the patient, often described as
passive cooperation. However, the operator must take considerable care with the test
procedure. Because of the freedom offered to patients by simply breathing normally,
moment-to-moment changes in respiratory resistance may be anticipated. Accordingly, a
minimum of three technically acceptable FOT tests of 20–30 s duration or longer should
be performed. The mouthpiece of the FOT instrument must be supported at a position to
ensure maintenance of a neutral relaxed head and neck posture, avoiding body postures
that might affect Zrs. Children should be seated in an appropriately-sized chair to
comfortably support their legs and adults should avoid crossing their legs, which requires
abdominal muscle contraction that may lead to end-expiratory lung volumes below
relaxation volume. Patients should be comfortably relaxed to maintain a constant body
84
OSCILLOMETRY: FOT AND IOS
position without muscular effort. In contrast, firm contraction of the facial muscles is
necessary to support airtight closure of the lips about the mouthpiece. It is also desirable
that FOT testing be performed in a quiet examination room, sufficiently far from others
undergoing spirometry so as to be undisturbed by vigorous operator instructions for
spirometry. The operator must review each FOT test immediately to ensure adequate
recording time free of artefacts; a minimum of 20 s of consecutive artefact-free recording
time is advisable. Lack of attention to these fundamental principles may result in highly
variable FOT tests in an individual that are not clinically interpretable. This is not a
troublesome issue in experienced clinical investigators’ laboratories but it is an important
consideration in clinical PFT laboratories who have not previously used FOT.
An important perceived concern about the clinical relevance of FOT is the availability
of a normal database from which to judge results in a particular patient. Published
reports of normative FOT data in children and adults are available, but the number and
size of these studies represent a much smaller normal population than is available for
spirometry [6, 32, 39, 74, 79, 80, 93]. Differences in techniques of FOT and more recently
in mouthpiece design may allow some degree of uncertainty concerning ranges of normal
expected values for both Rrs and Xrs, in both adults and children. Nevertheless, the
similarity of FOT data using monofrequency sine-wave oscillations, PRN or pulseshaped multifrequency FOT over the past four decades supports the acceptance of
clinically useful guidelines at this time. As expected, Rrs and Xrs are dependent on body
size, and recent data suggest the possibility of racial/ethnic differences [99].
It is suggested that concern over precise definition of "normality versus abnormality" in
an individual should not preclude clinical implementation of FOT at this time, because
bronchial and therapeutic challenges are very helpful in assessing airway responsivity. If
initial baseline Rrs/Xrs data do not clearly identify abnormality relative to existing
normative data, retesting after b-agonist inhalation will immediately identify increased
airway responsivity, and sequential studies over time provide similarly useful guidelines
for clinical patient management.
Diagnostic evaluations
The most obvious relationship with other PFT procedures concerns the widespread use
of spirometry. At the outset, it must be emphasised that spirometry measures maximal
forced respiratory efforts, while FOT measures quiet breathing. Accordingly, it is not
appropriate to demand that FOT and spirometric parameters be closely correlated as a
mandatory requirement for FOT to be considered valid. For example, children with
asthma most commonly manifest normal spirometry [100] with no spirometric response
to inhaled b-agonist [42], while they may manifest abnormal baseline FOT parameters
that are responsive to therapeutic challenge [101]. At the other end of the spectrum,
patients with advanced chronic obstructive pulmonary disease (COPD) commonly
manifest marked dynamic airway compression during spirometry with little spirometric
response to pharmacological treatment, but may often manifest significant FOT
responses [45]. For these reasons, use of spirometry to define severity of obstructive lung
disease or receiver-operating characteristic of true sensitivity and specificity of
oscillometry must no longer be considered the optimal standard.
In some patients, spirometry cannot provide optimal clinical information. Patients
with significant neuromuscular disease are unable to provide the motive force needed for
clear interpretation of spirometric results [97, 98]. Patients with lung allograft
transplantation have obvious thoracic wall limitations that preclude truly maximal
respiratory efforts until many months after surgery, during which time it is often not
possible to detect adverse events, such as infection or acute allograft rejection by
85
H.J. SMITH ET AL.
spirometry alone. It is common for forced vital capacity (FVC) and FEV1 to increase
over the first 18–26 weeks post-lung transplantation [102]. Despite this apparent
improvement in spirometric parameters, peripheral airway disease, if it occurs, will lead
to substantial worsening of FOT parameters [103] or of gas distribution [104].
Spirometric determination of responses to bronchial or therapeutic challenge may be
limited by the necessary deep inspiration immediately prior to maximal expiration,
allowing for distinctly different FOT responses [105–108]. This disagreement between
spirometric and FOT parameters during bronchial challenge may be readily documented
by IOS during quiet breathing immediately before and after a deep inspiration, which
commonly reveals immediate but transient bronchodilation in asthmatic subjects during
bronchial challenge. Figure 5 illustrates a 70-s recording of tidal breathing and a deep
inspiration with simultaneous display of calculated magnitude of impedance at 5 Hz
(Zrs5) using impulse oscillometry.
Rrs5 in this patient had increased markedly from 0.3 kPa?s?L-1 at baseline to
1.2 kPa?s?L-1 after cumulative inhalation of 0.25, 0.5 and 1.0 mg?mL-1 methacholine,
when FEV1 had changed by only 260 mL from a baseline of 2.73 L. Zrs5 during normal
breathing and following a deep inspiration with relaxed expiration shows that, at the
onset of the IOS test, Zrs5 is much increased, while after an inspiration to maximal lung
volume, there is a marked fall in Zrs5, which gradually "recovers" with time. Since deep
inspiration to maximal lung volume must immediately precede the FEV1 measurement,
patients like the one whose data are shown in figure 5 will manifest FEV1 responses to
methacholine that are altered by the immediate decrease in airway smooth muscle tone
following maximal inspiration [106, 107].
With these caveats it is suggested that FOT can provide useful supplementary
information to spirometry that may not be tightly correlated with spirometric results. To
put such FOT information into proper perspective, it is necessary to review commonly
observed patterns of FOT results in lung disease (primarily airflow obstruction), and in
response to bronchial and therapeutic challenge.
3
Volume L
2
1
0
-1
2.0
Zrs5 kPa·s·L-1
1.6
1.2
0.8
0.4
0.0
0
10
20
30
40
Recording time s
50
60
70
Fig. 5. – Tidal volume and magnitude of respiratory impedance at 5 Hz (Zrs5) plotted as a function of time
during methacholine challenge. First 40 s are resting breathing. After 40 s, subject inspired to total lung
capacity, followed by relaxation back to normal resting breathing. Note that Zrs5 increases markedly during
each exhalation. After the deep inspiration, Zrs5 decreases transiently, with gradual return towards initial levels
over the following 24 s. The moving average of Zrs5 is shown by the dash-dot line.
86
OSCILLOMETRY: FOT AND IOS
Oscillometry in relationship to other diagnostic pulmonary functional tests
An important body of work has related FOT to body plethysmography [29, 76–78]. A
group led by Van Noord have reported high correlations between Raw and FOT
parameters in patients with obstructive lung disease and between absolute total lung
capacity (TLC) and FOT parameters in patients with diffuse interstitial lung disease. In
further studies comparing plethysmography, spirometry and FOT to assess reversibility
of airflow obstruction, Van Noord’s group reported the distinctly lower sensitivity of
FOT than plethysmography [77]. The importance of this and other work by the Van
Noord group is discussed further in the following sections.
It is widely recognised that body plethysmographic resistance, Raw includes only the
resistance of the extrathoracic and intrathoracic airways, while Rrs includes that of the
chest wall and lung tissue in addition to airway resistance. Resistance of the chest wall
has been reported [75], but there has been limited clinical interest in this parameter
because of the technical difficulty of the measurements. Another difference between Raw
and Rrs relates to the status of the glottic aperature: it is commonly assumed that the
glottis is maximally open during panting, but Jackson et al. [109] have shown that this
occurs only in totally unrestricted panting. During voluntary attempts to control panting
frequency and tidal volume, there is significant adduction of the vocal cords. Similarly,
during quiet breathing, normal subjects commonly manifest a small, but variable, degree
of vocal cord adduction during expiration. In patients with obstructive lung disease, this
phasic expiratory adduction, visualised during the course of bronchscopy, does not
appear to be systematically different from that observed in normal subjects. Thus, it is to
be expected that average Rrs will differ systematically from panting plethysmographic
Raw.
It is also well established that Raw is more prominently influenced by large airway than
by small airway resistance. Thus, Smith and Dubois [110] reported a comparable increase
in deadspace when compared to the decrease in Raw in response to scopolamine in
normal subjects. In addition, Hensley et al. [111] reported similar changes in Raw and
deadspace after inhaled atropine. These results are consistent with the idea that Raw is
primarily influenced by large airways. In contrast, Rrs is importantly influenced by small
airway resistance, and, accordingly, it may be expected that FOT responses to
interventions that improve peripheral airway obstruction will be more prominent than
Raw responses. Because of the prominent effect of peripheral airway obstruction on FOT
measurements, it may be expected that FOT indices of peripheral airway obstruction
might correlate more closely with indices of gas distribution ("Closing volume" [104]) and
areas of lung hyperinflation manifested by computerised tomography [112] than with
plethysmographic or spirometric indices, although there are no published comparisons at
this time.
Oscillometry as a clinical monitor of response to treatment
By way of summarising the clinical relevance of FOT, it is worth considering the
special value of FOT as a means of monitoring response to interventions. FOT has been
reported to show greater sensitivity to inhaled corticosteroid or to b-agonist inhalation
[8, 113–115] than spirometry. Both inhaled corticosteroids and b-agonists improve small
airways function, and FOT responses manifest prominent changes in indices of
peripheral airway obstruction. In contrast, spirometric sensitivity to small airways
function is less prominent. Accordingly, it is expected that FOT might provide useful
indices of peripheral airway change in response to therapeutic interventions. Such use of
FOT provides a clinically valuable monitoring tool to follow therapeutic changes in small
87
H.J. SMITH ET AL.
airways function over time. This use of FOT for therapeutic monitoring is not dependent
on the use of FOT as an initial diagnostic evaluation.
Finally, as a matter of practical convenience, FOT is more readily utilised in the
clinical pulmonary function laboratory than body plethysmography. This issue is
relevant to recent interest in the therapeutic value of anticholinergics in patients with
COPD. As noted above, anticholinergics result primarily in large airway bronchodilation, and changes in Raw and deadspace are considerable [110, 111]. Thus, effective
airway cholinergic blockade decreases large airway bronchomotor tone and increases
deadspace, with relatively little effect on small peripheral airways disease. Whereas body
plethysmography may be considered a useful technique to monitor such treatment
effects, it is substantially less convenient to use routinely than FOT.
Oscillometry in the clinical pulmonary laboratory
Clinical interpretations of FOT responses in patients have often been related directly
to the application of a particular mechanical or electrical model of the respiratory
system. Van Noord et al. [78] discussed their results in diffuse interstitial lung disease
with respect to an electrical analogue of the respiratory system. They further confirmed
earlier work (vide infra) that ascribed negative frequency dependence of resistance to
peripheral airway obstruction [116, 117]. Engineering models of the respiratory system
have provided predictions of FOT characteristics in normal human subjects, and changes
in FOT parameters in lung disease. However, the fact that many of these predictions have
been observed empirically does not constitute proof of validity of one or other
engineering models. Rather, it provides evidence that under the particular conditions of
the FOT measurements undertaken, the empirical results show patterns that would be
intuitively expected in normal subjects and those with lung disease. Over the past 3
decades, a body of empirical evidence has accumulated that relates FOT results to
particular lung diseases, indeed establishing patterns that are characteristic of lung
disease. The following sections draw heavily upon this clinical research and codify FOT
results with respect to obstructive lung disease, with very limited data in diffuse
interstitial lung disease. These FOT data are not intended as validations of engineering
models, but instead, to illustrate commonly observed patterns of FOT characteristics
associated with lung disease.
Obstructive lung disease
The relationships of FOT to spirometry noted above have a common theme.
Spirometry does not provide a clear indication of peripheral airway obstruction, despite
the general information contained within the shape of the flow–volume curve, and values
of mid-flow rates (forced expiratory flow between 25 and 75% of the forced vital
capacity). Thus, the most striking characteristic of FOT in relation to spirometry is the
relatively greater sensitivity of FOT to peripheral airway disease [2, 18, 25, 29, 32, 42, 45,
52, 68, 82].
Peripheral airway disease. The most well-known FOT result empirically observed in
peripheral airway disease in adults is frequency dependence of resistance (fdr). Grimby
et al. [116], using multiple replicates of monofrequency FOT, were the first to demonstrate
the pattern of frequency dependence, wherein calculated Rrs was greater at 3 Hz than at 5,
7 or 9 Hz in patients with chronic airflow obstruction (CAO). As calculated Rrs decreased
as oscillation frequency increased, patients with CAO might manifest nearly normal
88
OSCILLOMETRY: FOT AND IOS
values of Rrs at sufficiently high frequencies. For this reason, Grimby et al. [116] focused
on low (3 Hz) monofrequency Rrs to avoid masking differences between patients with
airflow obstruction and normal subjects [118]. Many subsequent reports [18, 32, 36, 74,
82, 117, 119] confirmed that subjects with early peripheral airways disease, including
smokers, certain industrial workers and normal subjects after histamine infusion,
manifested frequency dependence, even with normal values of low-frequency Rrs in
smokers. Importantly, the abnormal frequency dependence of resistance occurred in the
presence of normal spirometry in those subjects with early peripheral airways disease [18,
32, 117]. This body of clinical research is largely empirical, although it had been shown on
autopsy many years earlier that cigarette smokers who died early in life had manifested
peripheral airway inflammation on autopsy [120]. Similarly, there is now ample evidence
of peripheral airway inflammation in patients with asthma, and, as will be illustrated
below, frequency dependence of resistance occurs prominently in asthma.
The sensitivity of frequency dependence to peripheral airway disease is the first
discriminant between methods of FOT in general: while monofrequency FOT is
convenient to dissect within-breath patterns of change in Rrs [61, 68] or changes in Rrs
during sleep-disordered breathing or in patients on mechanical ventilators [2],
multifrequency FOT is most convenient to document frequency dependence of resistance
in practical use in the clinical pulmonary function laboratory. Monofrequency FOT may
be used at two or more single frequencies; however, multifrequency FOT uses different
oscillation frequencies applied within a single burst to dissect patterns consistent with
peripheral rather than central airway obstruction. This dissection is based upon
established observations that pressure oscillations at frequencies w15 Hz are severely
damped out before reaching peripheral airways, while those at frequencies v10–15 Hz
penetrate much further to the lung periphery [25, 121].
The transition between "large central" airways and "small peripheral" airways is
neither precisely fixed anatomically nor precisely defined in terms of airway lumen
diameter. The illustrations in figures 6 and 7 reflect common patterns observed in
children with asthma and in adults with COPD.
Figure 6 shows representative IOS tests pre- and post-salbutamol in a 6-yr-old patient
Resistance Rrs kPa·s·L-1
1.0
0.8
0.6
0.4
0.2
0.0
0
5
10
20
15
Frequency ƒ Hz
25
30
35
Fig. 6. – Conventional plots of respiratory resistance (Rrs(f)), as a function of oscillation frequency, in a 6-yr-old
child with asthma. Note that frequency axis is shown between zero and 35 Hz, while data are plotted between
3–35 Hz. A vertical dotted line is shown at 5 Hz, the lower limit at which most impulse oscillometry system
data are reported. ––: data prior to nebulised b-agonist bronchodilator; ----: data after bronchodilator.
89
H.J. SMITH ET AL.
Resistance Rrs kPa·s·L-1
1.0
0.8
0.6
0.4
0.2
0.0
0
5
10
20
15
Frequency ƒ Hz
25
30
35
Fig. 7. – Respiratory resistance (Rrs(f)) in an adult patient with chronic obstructive pre- and post-nebulised bagonist bronchodilator. ––: pre-bronchodilator; ----: post-bronchodilator. Note that Rrs(f) is unchanged after
bronchodilator at frequencies w12 Hz.
with mild asthma. Rrs is plotted as a function of oscillation frequency (Rrs(f)-tracing).
Note that prior to b-agonist, Rrs5 is 0.93 kPa?s?L-1. Rrs falls steeply with increasing
oscillation frequency to a minimum at 18 Hz, after which it increases with further
increases in oscillation frequency. While Clement et al. [39] have shown that normal
children manifest a mild degree of frequency dependence, the very large fall in Rrs
between 5 and 15 Hz in figure 6 is consistent with abnormal peripheral airways function
in a 6 yr old. This is confirmed by administration of nebulised b-agonist, after which Rrs5
decreased to 0.59 kPa?s?L-1 (37% change). Note also that the fall of Rrs between 5–15 Hz
post b-agonist is much less than pre b-agonist. Baseline and post b-agonist IOS data in
this child may be compared with data in normal children of this age, who manifested
v15% fall in Rrs from 5–15 Hz [122]. The response to b-agonist in figure 6 may also be
compared with responses of normal nonatopic children who manifested an average
change in Rrs5 after salbutamol of 19% [122]. Finally, it can be seen in figure 6 that Rrs(f)
at frequencies w20 Hz decreased substantially after b-agonist. Adult patients with
asthma show similar findings to those in figure 6. Rrs may be nearly independent of
oscillation frequency in adult asthmatics after beta agonist.
Figure 7 illustrates an adult patient with COPD, pre- and post-b-agonist. In contrast
to asthma, Rrs(f) decreases continuously with oscillation frequency between 5–25 Hz at
baseline, and after b-agonist there is a decrease in Rrs(f) only at low frequencies,v12 Hz.
The findings in figures 6 and 7 are consistent with bronchodilation produced by bagonist in both large and small airways in the asthmatic subjects, but only in small
airways in the patient with COPD. In some patients with COPD who also have reactive
airways, b-agonist results in bronchodilation of large airways as well.
Central airway obstruction. Because of the frequency-dependent distribution of
oscillatory pressures within the airway tree, FOT provides separate, although not entirely
independent, indices of large and small airway responses. Thus, changes of resistance in
the larger airways are manifest by FOT as uniform changes in Rrs at all oscillation
frequencies, both low and high. An increase in resistance of the large airways was reported
in a recent study of rescue workers at the World Trade Center site in New York [18].
Rescue workers with no history of cigarette smoking exposed to large-particle air
90
OSCILLOMETRY: FOT AND IOS
pollution at the World Trade Center site manifested, at baseline, uniformly increased Rrs
at all oscillatory frequencies studied between 5–35 Hz, as shown in figure 8, pre- and postnebulised b-agonist.
Ironworkers with a history of cigarette smoking showed greater baseline increases in
low-frequency Rrs5, and, accordingly, a frequency dependence of resistance that is
characteristic of peripheral airway obstruction [7, 18, 29, 36, 82, 116, 117]. Responses to
nebulised b-agonist showed uniform decreases in Rrs across all frequencies in
nonsmokers, while smokers manifested larger decreases in Rrs5 than in Rrs20 [18].
Nonresistive components of forced oscillation technique. A second discriminant
between mono- and multifrequency FOT concerns reactive components of applied
pressure oscillations, reactance Xrs(f)-tracings, which appear prominently in
multifrequency FOT. Reactance can be assessed by computer-assisted analyses
utilising FFT; however, monofrequency FOT has not been widely adapted to
conveniently calculate Xrs.
Just as with the interpretation of Rrs, oscillation frequency provides a means of
examining different regions of the airway tree using Xrs (vide infra): at low oscillation
frequencies, elastic elements in peripheral airways are the dominant reactant to applied
pressure, and reflect small airway mechanical properties. In airflow obstruction, small
airways are functionally obstructed, due to peripheral airway inflammation in both
asthma and COPD. This results in portions of the distal lung periphery that are "in the
shadow" of effective obstruction of small airways. This pathological process leads to an
increase in the magnitude of Xrs at low frequencies. At high frequencies, accelerative
forces are the predominant reactant to applied pressures and occur virtually exclusively
in large airways where they are related to inertial properties. It should be noted that Van
Noord et al. [78] reported qualitatively similar changes in Xrs(f) tracings in patients with
diffuse interstitial lung disease and obstructive lung disease. Accordingly, changes in lowfrequency Xrs are not specific to obstructive lung disease, but rather reflect peripheral
airway disease. Mechanical interpretations of changes in low-frequency Xrs in obstructive
and restrictive lung disease are considered in detail in the methodology section.
Resistance Rrs kPa·s·L-1
1.0
0.8
0.6
0.4
0.2
0.0
0
5
10
20
15
Frequency ƒ Hz
25
30
35
Fig. 8. – Respiratory resistance (Rrs(f)), plotted as a function of oscillatory frequency pre- (––) and post (----)
-bronchodilator in an ironworker exposed to large-particle air pollution at the World Trade Center site. Note
increased Rrs at all frequencies, with no significant frequency dependence of resistance at baseline. After
nebulised b-agonist, Rrs(f) decreases in a parallel manner relative to baseline pre-bronchodilator.
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H.J. SMITH ET AL.
Figure 9 illustrates IOS Xrs(f) tracings in the asthmatic child whose Rrs(f) tracings preand post-nebulised b-agonist are shown in figure 6. Figure 9 shows that after b-agonist,
Xrs at 5 Hz (Xrs5) changes from -0.31 to -0.26 kPa?s?L-1 and the frequency at which Xrs is
zero (resonant frequency = fres, vide infra) changes from 18 to 17 Hz (6%). More
strikingly, the overall curvature of the Xrs(f) tracing changes from concave to the zero-X
axis to being convex to the zero-X axis. This change in curvature is consistent with Xrs(f)
tracings described by Clement et al. [36], who reported that patients with airflow
obstruction manifested a loss of the downward concavity of Xrs(f) tracings that is
commonly seen in normal subjects. b-agonist produced only small changes in Xrs5 and
fres in figure 9; however, the change in the overall Xrs(f) tracing curvature at frequencies
below fres was dramatic.
Previous investigations have emphasised that low-frequency Xrs and Rrs are most
sensitive to changes in peripheral airway function, and Xrs5 has been used as a primary
efficacy variable [113–115, 122, 123]. However, in small children, respiratory frequency is
commonlyw20–30 breaths?min-1, and, accordingly, the higher harmonics of fundamental
respiratory frequency may encroach on the lowest FOT frequencies analysed [41, 68]. As
a result, Xrs5 manifests relatively greater measurement noise. Some IOS studies in
children have reported failure of Xrs5 to manifest significant changes after inhaled
corticosteroids or b-agonists [42, 115, 124]. Marotta et al. [42] showed that the Xrs10
response to b-agonist, but not the Xrs5 response, manifested a significant difference
between asthmatic and nonasthmatic atopic children. This was associated with less
variability in Xrs10 responses compared with Xrs5.
The absolute value of Xrs (|Xrs|) changes differently as a function of oscillation
frequency below and above fres. At low frequencies of oscillation, below fres, |Xrs|
decreases as oscillation frequency increases up to fres. At fres, |Xrs| is zero. As oscillation
frequency increases above fres, |Xrs| increases with further increases in oscillation
frequency.
Because of this difference in the relationship between magnitude of Xrs and oscillation
frequency below and above fres, a quantitative index of Xrs magnitude at frequencies
below fres was developed by integrating all negative values of Xrs [18, 45, 52]. This index,
0.5
Reactance Xrs kPa·s·L-1
0.4
0.3
0.2
0.1
AX post
0
-0.1
-0.2
-0.3
AX pre
-0.4
-0.5
0
5
10
15
20
Frequency ƒ Hz
25
30
35
Fig. 9. – Respiratory reactance (Xrs(f)), plotted as a function of oscillatory frequency pre- (––) and post- (----)
bronchodilator. Same subject as figure 6. The integrated low-frequency reactance area (AX), is shown by vertical
hatching pre-bronchodilator, and by diagonal hatching post-bronchodilator. This area is reduced by y50% from
pre- to post-bronchodilator, associated with marked change in curvature of the Xrs(f) tracing. In comparison,
there are small changes in Xrs at 5 Hz and resonant frequency from pre- to post-bronchodilator.
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OSCILLOMETRY: FOT AND IOS
designated AX, provides an integrative function to include changes in the magnitude of
low-frequency Xrs, changes in fres and changes in curvature of the Xrs(f) tracing. AX
includes Xrs magnitudes at 5 Hz and slightly higher oscillation frequencies which
manifest improved signal-to-noise ratio, as noted above for atopic asthmatic children
[42]. It is represented graphically in figure 9 as the area under the zero Xrs axis above the
Xrs(f) tracing. As discussed below, this integrative index provides a single quantity that
reflects changes in the degree of peripheral airway obstruction and correlates closely with
frequency dependence of resistance. In figure 9, AX decreases by 50% after b-agonist,
closely comparable to the decrease in frequency dependence of resistance measured
between 5 and 15 Hz in the same child shown in figure 6, when R5-R15 decreased from
0.36 kPa?s?L-1 to 0.16 kPa?s?L-1 after b-agonist.
The perspective presented here of AX in relation to peripheral airway function, results
directly from the details presented in the methodology section, namely that lowfrequency Xrs essentially expresses the ability of the respiratory tract to store capacitive
energy, which is primarily resident in the lung periphery. In contrast, at frequencies
above fres, I, which is primarily resident in proximal conducting airways, contributes
significantly to the dissipation of externally applied pressures [6, 46]. Thus, oscillation
frequencies above fres reflect mechanical properties of more proximal conducting
airways. Because of these differences in mechanical properties reflected by low- and highfrequency oscillation, calculation of the arithmetic mean Xrs is not likely to be optimally
useful to assess pulmonary mechanical responses. Thus, Van Noord et al. [29] reported
that the mean value of Xrs change was less sensitive than FEV1 in assessing the effect of
histamine. However, their graphic mean Xrs–frequency data reveal changes in the
estimate of AX from y0.15 at baseline, to 1.7 or 2.5 kPa?L-1 when mean decrease in
specific airway conductance (sGaw) was 40% or 15% in FEV1. These increases in AX of
1,000–1,500% are comparable to those measured using IOS in the current author’s
laboratory during methacholine challenge: the patient shown in figure 6 manifested an
increase in AX from 0.34 at baseline to 7.5 kPa?L-1 (w2,000% increase) after cumulative
exposure to methacholine up to 1.0 mg?mL-1. Furthermore, the study of Van Noord
et al. [77] during assessment of reversibility of airflow obstruction by FOT, body
plethysmography and spirometry reported that changes in mean Xrs did not contribute
significantly to discriminant function beyond spirometry, Raw and Rrs at 6 Hz. In
contrast, estimated AX, approximated from their graphic mean Xrs–frequency data,
showed a 50% reduction after salbutamol, from y3.1 to 1.5 kPa?L-1. Both the baseline
AX in patients with airflow obstruction and the AX decreases in response to b-agonist are
comparable to those recently reported using IOS [101].
The empirical observations discussed above do not prove that any particular
engineering model is a true representation of the lung, especially the diseased lung as
emphasised by Van Noord et al. [78]. Nevertheless, FOT results predicted by models
may correlate usefully with independent clinical physiological and pathophysiological
evidence. Quite apart from engineering models, an intuitive understanding of Xrs may be
appreciated from the physical principles elucidated above and amplified in the foregoing
discussion of methodology. The applicability of high frequencies to large central airways
and low frequencies to peripheral airways is not a consequence of any particular
engineering model, but is observed empirically both in Rrs and Xrs.
Clinical interpretation of forced oscillation technique
As also noted in the methodology discussion of FOT, abnormalities of Xrs are not
specific to obstructive lung disease, because these same patterns have been reported in
lung fibrosis [78]. In clinical diagnostic lung function testing including FOT, spirometry,
93
H.J. SMITH ET AL.
gas diffusion and body plethysmography, the problem is not usually distinguishing
between obstructive and restrictive disease. The more important issue is the relative
severity of pathophysiological abnormality. Furthermore, there is evidence to suggest
that the early pathological lesion in lung fibrosis is inflammation of the small airways
[125]. Thus, changes in Xrs magnitude in lung fibrosis and in peripheral airflow
obstruction may both reflect peripheral airway inflammation. If these nonspecific Xrs
abnormalities in lung fibrosis represent small airway inflammation, they may respond to
anti-inflammatory treatment, analogous to the way asthmatic peripheral airway
inflammation responds to corticosteroids. Such responses are more likely to be found
in FOT parameters than in spirometry. Clinical interpretation may then be considered in
the setting of response to treatment interventions.
Clinical interpretation of FOT can be related both to effects of airway smooth muscle
tone and airway inflammation. Below, effects of anticholinergic, b-agonist or
corticosteroid medications are represented, because these agents are most commonly
utilised clinically. Commonly observed changes in FOT measures in patients with
obstructive lung disease are illustrated. Rrs is considered first.
Rrs effects. While inflammation is a cellular process, it has mechanical consequences.
These consequences may be considered in relation to bronchoconstriction, defined as
increased tone of airway smooth muscles, and the common perception of bronchodilation
defined as a decrease in smooth muscle tone.
When airway smooth muscle tone increases, Rrs increases because of decreased airway
lumen. Airway lumen is also decreased with inflammation or oedema in the walls of the
airways. Therefore, Rrs increases as a result of inflammation and oedema. Peripheral
airways have much smaller lumina than central (large) airways, and inflammation/
oedema in the walls of peripheral airways can reasonably be expected to have a
proportionately larger effect on lumen size than inflammation/oedema in larger airways.
In the discussion that follows, "low-frequency Rrs" will be denoted as Rrs at frequencies
v15 Hz and "high-frequency Rrs" as Rrs at frequencies w20 Hz, with the latter term
synonymous with large airway resistance.
If an intervention, such as inhaled anticholinergic, achieves bronchodilation with no
effect on inflammation, it may be expected that large airway lumen will increase, with
little effect on peripheral airway lumen. In this event, large airway Rrs will decrease. Lowfrequency Rrs will decrease to a similar degree and little or no change in frequency
dependence occurs. This is illustrated in a 55-yr-old patient with COPD in figure 10.
If an intervention such as inhaled b-agonist achieves bronchodilation with little or no
effect on inflammation, it may be expected that peripheral airway lumina will increase, in
addition to any release of bronchoconstriction in large airways. In this case, lowfrequency Rrs will decrease out of proportion to high-frequency Rrs.
In asthmatic patients, both high-frequency and low-frequency Rrs may decrease, with
relatively greater decrease in low-frequency Rrs and associated decrease in frequency
dependence of resistance as illustrated in figure 6.
In patients with COPD, a decrease in low-frequency Rrs after b-agonist with little or
no change in high-frequency Rrs is commonly observed, as illustrated in figure 7. In COPD
patients with lung hyperinflation, little or no decrease in Rrs may occur after
b-agonist inhalation, particularly if there is an associated fall in resting end-expiratory
lung volume. If Rrs remains the same after b-agonist while end-expiratory lung volume
decreases, this represents "functional bronchodilation", because the same resistance
pertains at lower operating lung volumes. Accordingly, failure of Rrs to decrease in patients
with COPD need not be considered as "no response" to b-agonist bronchodilator.
If an intervention such as inhaled corticosteroids achieves a decrease in inflammation
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OSCILLOMETRY: FOT AND IOS
Resistance Rrs kPa·s·L-1
2.0
1.6
1.2
0.8
0.4
0.0
0
5
10
20
15
Frequency ƒ Hz
25
30
35
Fig. 10. – Respiratory resistance (Rrs), plotted as a function of oscillation frequency in a 55-yr-old male with
chronic obstructive pulmonary disease pre- (––) and post- (----) anticholinergic bronchodilator. Note that Rrs at
5 Hz is markedly elevated with marked frequency dependence of Rrs. After inhaled anticholinergic bronchodilator, there is a significant decrease in Rrs that is nearly identical at all frequencies, indicating a decrease in
proximal airway resistance, with little or no effect on peripheral airways
with no effect on airway smooth muscle tone, FOT responses might be expected to reflect
a relatively greater impact due to decrease in peripheral airway inflammation with
resultant increase in peripheral airway lumina. Such an effect results in a significant
"dilation" of peripheral airways due to decreased inflammatory encroachment on
peripheral airway lumina. Thus, a decrease in peripheral airway resistance can be
expected, manifest as a greater decrease in low-frequency than in high-frequency Rrs, and
concomitant decrease in frequency dependence.
In asthmatic patients, a decrease in large airway Rrs (high-frequency Rrs) may also
occur. In patients with COPD, there may be a decrease in low-frequency Rrs; however,
little or no decrease in Rrs may be manifest, especially if lung hyperinflation is present.
Xrs effects. How then does decreased inflammation manifest itself in patients with
COPD? As noted in the preceding section, describing nonresistive components of FOT,
the magnitude of low-frequency Xrs is increased in COPD due to functional peripheral
airway obstruction, with resultant contraction of surface area of the lung periphery
exposed to low-frequency oscillations. Indeed, low-frequency Xrs is more sensitive to
peripheral airway obstruction in COPD/emphysema than Rrs. In the presence of
peripheral airway obstruction in patients with COPD, relatively small increments to
airways resistance may occur because of the large cumulative cross-sectional diameter of
all airways in the lower generations of airways, as manifest in the trumpet model of
Weibel [126]. Accordingly, body plethysmographic measurement of airway resistance
may be nearly normal, and only by measuring absolute thoracic gas volume is the
abnormality manifest.
If peripheral airway inflammation is decreased by administration of inhaled
corticosteroids, peripheral airway lumina increase and the patency of small airways
expands in the direction of the lung periphery. As a result, a portion of the lung periphery
comes "out of the shadow" of small airway obstruction and a larger surface area is
presented to the low-frequency oscillations. This acts to decrease the magnitude of lowfrequency Xrs. This is illustrated in figure 11a showing IOS tracings in a COPD patient at
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H.J. SMITH ET AL.
a) 0.5
Reactance Xrs kPa·s·L-1
0.4
0.3
0.2
0.1
AX post
0.0
-0.1
-0.2
-0.3
AX pre
-0.4
-0.5
Resistance Rrs kPa·s·L-1
b)
1.0
0.8
0.6
0.4
0.2
0.0
0
5
10
15
20
Frequency ƒ Hz
25
30
35
Fig. 11. – a) Respiratory reactance (Xrs), plotted as a function of oscillatory frequency in a 73-yr-old male with
chronic obstructive pulmonary disease before and after 4 weeks inhaled corticosteroid (ICS) therapy. ––: preICS; ----: post-ICS. Integrated low-frequency reactance area (AX) is vertically hatched pre-ICS and diagonally
hatched post-ICS. AX is decreased by y50% after 4 weeks of therapy, resonant frequency by 10% and Xrs at
5 Hz decreased by 0.12 kPa?s?L-1. b) Respiratory resistance (Rrs) plotted as a function of oscillation frequency in
the same patient. ––: pre-ICS; ----: post-ICS. Note that Rrs at 20–22 Hz is relatively unchanged, while Rrs at 5
and 10 Hz are substantially decreased. High-frequency resistance (Rrs at 25–35 Hz) is decreased after ICS
therapy. See text for discussion.
baseline (–––) and after 4 weeks of inhaled corticosteroids treatment (-----). AX decreased
from 3.0 to 1.3 kPa?L-1 after inhaled corticosteroid treatment. In this patient, Rrs and
frequency dependence of the Rrs(f) tracing show significant decreases, as illustrated in
figure 11b; Rrs5 improved from 0.68 to 0.48 kPa?s?L-1, Rrs15 from 0.41 to 0.35 kPa?s?L-1.
Furthermore, this patient with COPD manifests somewhat "responding" large airways,
as his high-frequency Rrs (at 30–35 Hz) also decreased by y20%.
Figure 12 shows tracings in a COPD patient when stable at baseline and during the
onset of an exacerbation. Figure 12a shows baseline reactance area AX = 1.2 kPa?L-1,
with increases to 1.6, 2.4 and 4.3 kPa?L-1 over a duration of 9 days of exacerbation. Rrs5
increased significantly, but less dramatically from 0.51 at baseline to 0.54, 0.59 and
0.65 kPa?s?L-1 during exacerbation, shown in figure 12b. Frequency dependence,
calculated as the fall in Rrs from 5 to 15 Hz, was 0.15 kPa?s?L-1 when stable at baseline,
and increased to 0.2, 0.23 and 0.31 kPa?s?L-1 during exacerbation.
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OSCILLOMETRY: FOT AND IOS
a) 0.5
Reactance Xrs kPa·s·L-1
0.4
0.3
0.2
0.1
0.0
-0.1
-0.2
-0.3
-0.4
-0.5
Resistance Rrs kPa·s·L-1
b)
1.0
0.8
0.6
0.4
0.2
0.0
0
5
10
15
20
Frequency ƒ Hz
25
30
35
Fig. 12. – a) Respiratory reactance (Xrs(f)) in a 73-yr-old male chronic obstructive pulmonary disease patient
when stable at baseline and during onset of exacerbation over a duration of 9 days. Note progressive increase in
the reactance area (AX) from trial 1 to trial 4. b) Respiratory resistance (Rrs(f)) in the same patient. Note
progressive increase in Rrs at 5 Hz (Rrs5) from trial 1 to trial 4. Changes in Rrs5 were relatively smaller than
changes in AX over a duration of 8 days. –––: trial 1, baseline; -----: trial 2, 7 days after trial 1; –– - ––: trial 3,
8 days after trial 1; –– -- ––: trial 4, 9 days after trial 1.
The close correlation between changes in AX and changes in frequency dependence of
resistance in individual patients shown in figures 6–12 are confirmed across individuals in
the occupational study reported by Skloot et al. [18], who found a correlation of 0.92
between frequency dependence of resistance and AX across a sample of ironworkers at
the World Trade Center site, with variable exposure to cigarette smoking and largeparticle air pollution, and resultant variability in both large and small airway
obstruction. The close correlation between frequency dependence of resistance and
AX is consistent with both indices reflecting small airway function.
Coherence. Coherence, first introduced by Michaelson et al. [5], is defined as the autoand cross-correlations of phase and amplitude of oscillatory pressure and flow
components. It reflects, in an "engineering sense", the linearity of the respiratory
system and, in a "biological sense", the variability of the respiratory system from time to
time within the sample of data. Marked temporal variability of the respiratory system
within a breath occurs commonly in COPD, where even during quiet breathing, dynamic
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H.J. SMITH ET AL.
compression of intrathoracic airways may occur. As a result, Rrs and Xrs may increase
substantially during expiration.
Figure 13a illustrates a patient with severe COPD, with volume and Zrs5 as functions
of time. Figure 13b shows coherence plotted as a function of oscillation frequency, with
separate tracings for average combined inspiration/expiration as well as for inspirationonly and expiration-only.
Figure 13a shows marked changes in Zrs5 with respiratory phase, similar to those
shown by Marchal and Loos [68]. There is a clear decrease in Zrs5 at the onset of
inspiration, keeping minimal values until end-inspiration. A marked abrupt rise in Zrs5
occurs at the onset of expiration with elevated values including end-expiration. In figure
13b, the value of averaged coherence at 5 Hz, 0.4, is distinctly lower than at frequencies
i10 Hz.
Volume L
a) 1.5
1.0
0.5
0.0
Zrs5 kPa·s·L-1
-0.5
3
2
1
0
0
10
20
Time s
30
40
b) 1.0
0.9
Coherence g2
0.8
0.7
0.6
0.5
0.4
0.3
0.2
0.1
0.0
0
5
10
15
20
Frequency ƒ Hz
25
30
35
Fig. 13. – a) Volume and magnitude of respiratory impedance at 5 Hz (Zrs5) plotted as a function of time
during a 40-s impulse oscillometry system (IOS) test in a 53-yr-old patient with severe chronic obstructive
pulmonary disease. Note marked increase in Zrs5 during the expiratory phase of every tidal volume. b)
Coherence of all IOS data (global average (–––) and separated inspiratory (–– - ––) and expiratory (-----)
respiratory phases) plotted as a function of oscillation frequency during the 40-s test. Note low coherence for
average data (v0.5) at 5 Hz with prominent increase to 0.7 at 10 Hz. See text for discussion.
98
OSCILLOMETRY: FOT AND IOS
Numerical calculations of coherence during the separate respiratory phases reveal that
inspiratory-only and expiratory-only coherences are systematically greater than the
combined coherence over the entire spectrum of oscillatory frequencies. The tracings in
figure 13 are consistent with more uniformity of respiratory mechanics within the
separate inspiratory and expiratory phases than for the combined total breath average.
At 10 Hz, the coherence averaged across both respiratory phases is 0.7, while that for
separate inspiratory and expiratory data are both 0.9, reflecting differences in respiratory
mechanical parameters that pertain to the separate respiratory phases. Despite the very
low coherence for combined inspiration/expiration, three consecutive 40-s impulse
oscillometry system recordings obtained within 5 min manifested an average respiratory
resistance at 5 Hz of 1.08, 1.0 and 1.02 kPa?s?L-1 and reactance area of 5.9, 6.1 and
6.2 kPa?L-1. Thus, the within-phase uniformity of mechanical parameters reflected by
separate inspiratory and expiratory coherences is borne out by standard deviations of
v5% for triplicate measures.
Summary
The aim of this chapter has been to describe the unique and clinically relevant
information that forced oscillation technique (FOT) provides. This may be derived
without mathematical mastery of technological principles of the equipment and/or of
numerical models. It is emphasised that recognition of the change in respiratory
mechanical parameters as a function of oscillation frequency is necessary to appreciate
the outstanding value of FOT in its ability to assess peripheral airway function. This
has been one of the major challenges in respiratory diagnostics up to the present time.
The short duration of the FOT test, 20–30 s, makes it particularly useful as part of a
diagnostic programme of lung function testing; it is not suggested that FOT be used in
lieu of conventional pulmonary function testing, but rather in addition. FOT measures
resting breathing while spirometry assesses maximal respiratory performance of the
patient. The special value of FOT in terms of short-term response to bronchial and
therapeutic challenge has been emphasised as well as its value in monitoring long-term
trend responses to therapy.
The simplicity of FOT measurements and its minimal requirements on subject
cooperation are in rather sharp contrast to its current limited clinical acceptance. Two
primary reasons for the present limited application of FOT include the need for
viewing respiratory mechanical parameters over a range of frequencies and the
resultant central-peripheral specificity of oscillatory parameters, with specific
emphasis on the reflection of peripheral airway function by low-frequency reactance.
Indeed, lack of awareness of this ability of FOT to assess peripheral airway function
has turned physicians to the use of multiple replicates of high-resolution computed
tomography lung scans to assess small airway function. Other reasons for limited use
of FOT currently may include the greater variability of FOT measures compared with
spirometry. Despite such variability, use of at least three replicate FOT measures
combined with therapeutic challenge can provide sensitive evaluation of small airway
function.
The freedom allowed to the subject to breathe "naturally" imposes increased demands
for vigilance on the operator, who must maintain a quiet environment for forced
oscillation technique testing. Operators must also reassure subjects that their
relaxation is needed, except for the facial musculature ensuring tight lip closure on
the mouthpiece. Posture must be supported to maintain subject comfort and the
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instrument mouthpiece must be brought to the subject to avoid stretching of the neck.
Finally, the availability of results from a brief test must not lead the operator to accept
a single measurement, but rather, the usual clinical testing procedure of at least three
replicate measures is required.
Keywords: Forced oscillation technique, impulse oscillation technique, reactance,
resistance, respiratory impedance.
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105
CHAPTER 6
Pulmonary gas exchange
J.M.B. Hughes
National Heart and Lung Institute, Imperial College Faculty of Medicine, Hammersmith Hospital, London,
UK.
Correspondence: J.M.B. Hughes, 4 Cedars Road, London SW13 0HP, UK.
In this "overview", pulmonary gas exchange is considered in three sections. The first
section (Normal values, Causes of hypoxaemia I, Respiratory failure) contains "basic"
knowledge about pulmonary gas exchange, which is relevant to all who work in clinical
medicine. Normal values for arterial oxygen partial pressure, content and Hb saturation
(Pa,O2, Ca,O2, Sa,O2) are reviewed, and ventilation–perfusion mismatch and alveolar
hypoventilation are highlighted as the two commonest causes of a low Pa,O2 and
respiratory failure. Different types of respiratory failure are discussed, with special
emphasis on the hypoxaemia and hypercapnia occurring in chronic obstructive
pulmonary disease (COPD) patients in failure (alveolar ventilation–perfusion (V9A/Q9)
mismatch effect) and the rise in arterial carbon dioxide partial pressure (Pa,CO2) with
uncontrolled O2 therapy (alveolar hypoventilation effect).
The next section (Oxygen carriage in blood, Heterogeneity of ventilation and
perfusion, Causes of hypoxaemia II) focuses on "intermediate" knowledge, with which all
respiratory specialists should be familiar. First, the relationship between oxygen content
(CO2) (and oxygen Hb saturation (SO2)) and oxygen partial pressure (PO2), the so-called
oxygen dissociation curve (ODC), is introduced. The P50 (partial pressure at half
maximum blood concentration) for oxygen is defined. Heterogeneity of V9A and Q9
(leading to V9A/Q9 mismatch) is analysed using the PO2–PCO2 diagram, the Riley three
compartment model (physiological shunt flow/total pulmonary blood flow [Q9s/Q9T] and
dead space (dead space/tidal volume [VD/VT])), and the ideal alveolar–arterial PO2
gradient. The causes of hypoxaemia are outlined, and a possible overlap between
intrapulmonary shunt and diffusion limitation is discussed using the hepatopulmonary
syndrome as an example. The theoretical basis of transcutaneous measurements of Pa,O2
and Pa,CO2 (high skin blood flow and a narrow arteriovenous partial pressure difference)
and Sa,O2 is mentioned.
The last section (Oxygen affinity in special situations, Diffusion, Inert gas transport)
contains "advanced" knowledge, appropriate for staff in intensive care units,
anaethesiology or rehabilitation, or those undertaking research. The P50 for oxygen is
reviewed in special situations (right shift in exercise and some haemoglobinopathies, left
shift in CO poisoning, in utero and on the Mt Everest summit). The importance of gas
phase diffusion within the acinus is emphasised. The pathogenesis of alveolar–capillary
diffusion and diffusion limitation for oxygen on exercise (in lung fibrosis and at extreme
altitude) is explained. Lastly, inert gas transport is reviewed, focusing on the multiple
inert gas elimination technique (MIGET), a sophisticated analysis of V9A/Q9 mismatch,
which has provided information on the pathogenesis of hypoxaemia in different clinical
situations.
Eur Respir Mon, 2005, 31, 106–126. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
106
GAS EXCHANGE PRINCIPLES
Normal values
Arterial oxygen tension
Pa,O2 in normal subjects is affected by several factors: age, body mass index (BMI),
posture, altitude and inspired oxygen fraction (FI,O2: normal=0.21 or 21%).
The units of Pa,O2, Pa,CO2 are kilopascals (kPa) in Europe, but mmHg in North America
(1 kPa=7.5 mmHg).
Pa,O2 (kPa)~19:15{(0:052|age){(0:075|BMI){(0:076|Pa,CO2 ) ½SEE 1:0
ð1Þ
These values [1] were established in lifelong nonsmoking subjects with normal
pulmonary function. Pa,O2 on average declines by 0.55 kPa per 10 yrs from 13.3 kPa
at age 20 yrs to 10.7 kPa at age 70 yrs. Pa,O2 rises by about 1.3 kPa in pregnancy
(with a corresponding fall in Pa,CO2) [2], but there are no other sex effects. The fall in
Pa,O2 with age is caused by an increase in V9A/Q9 mismatching. In obese middle aged
and elderly subjects, Pa,O2 is lower in the supine position [3] due to dependent zone
bronchiolar collapse (and possibly atelectasis). People living or climbing at altitude,
or flying in pressurised aircraft at 9,850–10,770 m have a reduced partial pressure of
inspired oxygen (PI,O2) and, as a result, a low Pa,O2 [4]. They may compensate, to some
extent, by hyperventilating, which lowers Pa,CO2 and raises Pa,O2 (by a roughly
equivalent amount). The FI,O2 of air (y21%) is unchanged at altitude, but total
barometric pressure and the partial pressures of O2 and N2 fall.
Clinically, FI,O2 is often increased as a therapeutic measure; if 100% O2 is breathed at
sea level, Pa,O2 in normal subjects may rise from 13.3 to 80 kPa. It is important to know
the inspired concentration of oxygen when interpreting a Pa,O2 value.
The Pa,CO2 is not affected by age, but it is lowered by hyperventilation, the usual causes
being hypoxaemia, metabolic acidosis (e.g. diabetic) and anxiety.
Arterial oxygen content and saturation
Breathing air, 98.5% of oxygen in arterial blood is bound to haemoglobin. At a
"normal" Hb concentration (say 14.8 g?dL-1) each litre of arterial blood carries 200 mL
of oxygen, but only 100 mL?L-1 if [Hb] is 7.4 g?dL-1. At rest, the arteriovenous (a–v)
oxygen content difference is 50 mL?L-1 so that mixed venous blood (assuming a normal
cardiac output) contains 150 mL?L-1 (75% of the arterial value), but only 50 mL?L-1
(25% of the normal arterial content) at [Hb] 7.4 g?dL-1.
Ca,O2 is not measured routinely. It can be calculated (in mL?L-1) by plotting Pa,O2 on a
standard oxygen dissociation curve (ODC) (see section Oxygen carriage in blood and fig. 1)
reading off the percentage saturation of Hb (HbO2) with oxygen at that Pa,O2 and then
multiplying by the [Hb] and the O2 capacity of blood (1.39 g?dL-1). Nevertheless, it is now
very easy to measure HbO2 % saturation with a pulse oximeter attached to a finger or an ear
lobe. Pulse oximetry detects light transmitted at two wavelengths, corresponding to
deoxygenated and oxygenated haemoglobin. The signal is the difference in absorbance
between the peripheral systolic pulse wave and the subsequent diastole, a difference of
only 1–10% of the total light absorbance. Carboxyhaemoglobin (HbCO) (and methaemoglobin) absorb light at the same wavelength as deoxyhaemoglobin, so that HbO2 %
is overestimated in the presence of HbCO.
The requirements and reservations of pulse oximetry (Sp,O2) are shown in table 1. With
these reservations, pulse oximetry is acceptably accurate at rest and on exercise when
compared with simultaneous estimates of Sa,O2 from arterial blood samples (v2%
difference between estimates) [5]. The weakness of pulse oximetry is that it is insensitive
107
J.M.B. HUGHES
Mixed venous
50
P50
0
0
20
100
[93]
[89]
75
T°·PCO2·H+·2.3-DPG
50
Normal
range
Lower limit of normal
pH 7.2
P50
l
l
l
pH 7.6
pH 7.4
100
12 Arterial 16
l
Threshold for home oxygen
150
kPa
8 [9.3]
25
Dissolved O2 (plasma)
40
60 [70] 80
100
Oxygen partial pressure mmHg
120
Oxygen saturation (Sa,O2) %
T°·PCO2·H+·2.3-DPG
Urgent oxygen therapy required
Oxygen concentration mL·L-1
200
4 [5.3]
0
0
Fig. 1. – Oxygen concentration in blood (left axis) or per cent saturation of haemoglobin with oxygen (Sa,O2,
right axis) plotted against oxygen partial pressure (mmHg, bottom axis, kilopascals top axis) showing right and
left shifts of the curve produced by physiological variables. PCO2: carbon dioxide tension; P50: partial pressure at
half maximum blood concentration for oxygen; DPG: 2, 3-diphosphoglycerate.
to minor degrees of hypoxaemia; in the Pa,O2 range 13.3 down to 10 kPa HbO2 changes
only 3% (97.5 to 94.5%) because of the shape of the ODC (see fig. 1). The strength of
pulse oximetry is its ability to follow changes – from rest to exercise, from air to oxygen
breathing, and for continuous overnight monitoring. The laboratory and domiciliary
uses of pulse oximetry are shown below:
.
.
.
.
.
.
Home oxygen therapy assessment
Monitoring during exercise tests
Overnight monitoring for obstructive sleep apnoea (OSA) diagnosis
Monitoring at home (done by the patient) by day or at night
Assessment of "fitness to fly" using 15% inspired oxygen
Substitute for arterial sampling in children or for serial observations
Causes of hypoxaemia (low arterial oxygen partial pressure) I
Ventilation–perfusion mismatching
In the case of intrapulmonary disease, V9A/Q9 mismatching is nearly always the cause
of arterial hypoxaemia. To understand this, consider the case of suddenly (and
simultaneously) blocking the left main pulmonary artery (with an embolus) and the right
main bronchus (with a tumour that has bled). Without ventilation, the blood flow (equal
to the whole cardiac output) through the right lung would be unoxygenated (once the
small oxygen stores in the lung had been exhausted); the V9A/Q9 ratio would be zero and
108
GAS EXCHANGE PRINCIPLES
Table 1. – The reservations and requirements of pulse oximetry (Sp,O2)
Reservations
Requirements
Adequate arterial pulsation
Carboxy Hb v3%
Steady state
Skin pigmentation
Use vasodilator cream
Avoid smoking for 24 h
Wait for 5 min (minimum)
Not a problem, but avoid nail polish and very bright lighting
the Pa,O2 and Pa,CO2 of the blood leaving the right lung would have the same composition
as the mixed venous blood entering it. The other lung with ventilation but no blood flow
would act as a dead space with a V9A/Q9 ratio of infinity and an alveolar PO2 and PCO2
equal to that in inspired air. A mixture of V9A/Q9 ratios of zero and infinity means no
effective gas exchange. As V9A/Q9 ratios increase from zero and decrease from infinity,
gas exchange efficiency increases until the optimum ratio (y0.86) is reached. In real life,
there is a spread of V9A/Q9 values throughout the lung on either side of this "optimum"
value. The larger the spread, the greater the inefficiency of gas exchange. Low V9A/Q9
units lead to arterial hypoxaemia (and hypercapnia). High V9A/Q9 units contribute
wasted ventilation or "dead space". For further information, consult the PO2–PCO2
diagram discussed in the section Heterogeneity of ventilation and perfusion (see below).
The effect of V9A/Q9 mismatch raises Pa,CO2 as well as lowering Pa,O2, but the effect on
Pa,O2 is greater. In simple terms, this is because the "a–v" difference for PO2 (13.3–5.3=
8 kPa) is much greater than the "v–a" PCO2 difference (0.8 kPa). The body’s compensation
for hypoxaemia and hypercapnia is to increase minute ventilation (hyperventilation). If, for
example, V9A/Q9 mismatch has caused Pa,O2 to fall to 8 kPa [D5.33 kPa from "normal"] and
Pa,CO2 to rise to 6.8 kPa [D1.5 kPa], hyperventilation sufficient to cause a 2 kPa
improvement in both blood gas values will result in Pa,O2 10 kPa (still abnormal) and Pa,CO2
4.8 kPa (slightly low). With V9A/Q9 mismatch, Pa,O2 is nearly always reduced, but Pa,CO2
may be raised, normal or low depending on the ventilatory response. In the
"emphysematous" type of COPD, or "pink puffer", Pa,O2 may be surprisingly well
preserved (e.g. w11 kPa) but at the expense of hyperventilation and a low Pa,CO2.
Alveolar hypoventilation
An inadequate level of ventilation is the other main cause (in y5% of cases) of
hypoxaemia; its origin is usually extrapulmonary and the Pa,CO2 is always raised. It is
caused by insufficient alveolar ventilation [V9A] (total (V9E) minus anatomic dead space
(V9D) ventilation) in relation to metabolic demands of oxygen consumption (V9O2) and
carbon dioxide production (V9CO2). Respiratory centre depression (from anaesthetic,
sedative or analgesic drugs) or diseases affecting the diaphragm or its nerve supply, or
gross restriction of the chest wall (such as severe kyphoscoliosis) all lead to shallow
breathing, low V9E and inadequate V9A. Shallow breathing, in the long term, may lead to
retention of secretions and atelectasis (deep breaths assist in the renewal of the alveolar
surfactant lining). Oxygen breathing in exacerbations of COPD may lead to shallower
breathing and a further rise in Pa,CO2 (see section on Respiratory failure). The
consequence of alveolar hypoventilation for arterial blood gases is that Pa,O2 falls and
Pa,CO2 rises in roughly equal amounts. In theory, DPa,CO2/DPa,O2=0.8 (where 0.8 is the
respiratory quotient (RQ) imposed on the lung by body metabolism), but because of
accompanying V9A/Q9 mismatch, the fall in Pa,O2 may equal or exceed the rise in Pa,CO2.
Recognition of hypoventilation must take the clinical context into account rather than
relying on the Pa,O2–Pa,CO2 pattern, though a rise in Pa,CO2 is mandatory.
109
J.M.B. HUGHES
Respiratory failure
There is no precise definition of respiratory failure in terms of Pa,O2 and Pa,CO2. In
clinical terms, acute respiratory failure is an unstable condition when Pa,O2 and Pa,CO2 are
progressively falling and rising respectively. Chronic respiratory failure is a stable
condition associated with: 1) severe hypoxaemia (e.g. Pa,O2 v8 kPa, Sa,O2 v90%) without
(Type I) or with (Type II) hypercapnia; or 2) severe hypercapnia (Pa,CO2 w7 kPa) with
mild hypoxaemia (Pa,O2 w10 kPa) – the latter occurs with extrapulmonary conditions
associated with hypoventilation. The actual Pa,O2 and Pa,CO2 values defining "failure" are
somewhat arbitrary. The common causes of respiratory failure are shown in table 2.
The pathophysiology of different types of acute or acute-on-chronic respiratory failure
are set out in table 3. In acute pulmonary gas exchange failure (ARDS), with severe
V9A/Q9 mismatch and many gas exchange units flooded with plasma transudate or
exudates, corresponding to a V9A/Q9 of zero, i.e. "physiological shunt", severe hypoxaemia is the problem (Pa,O2 v5 kPa) and a high FI,O2 (w60%) may be required to achieve a
"safe" Pa,O2 level (w8 kPa). Central CO2 sensitivity remains normal, so hypercapnia does
not occur, if hyperventilation can be sustained. Nevertheless, such severe hypoxaemia
cannot be tolerated for long, and intermittent positive pressure ventilation (IPPV) with
positive end-expiratory pressure will be required. In extrapulmonary failure – whether
originating in the brain stem, phrenic nerves or the diaphragm itself – the weak link is not
pulmonary gas exchange, but the ability of brain, nerves or muscle to respond to the
hypercapnic stimulus. Since gas exchange is nearly normal, FI,O2 needs to be increased
only slightly, if at all. Ventilatory assistance with nasal intermittent positive pressure
(NIPPV), particularly at night, is the cornerstone of treatment.
Chronic obstructive pulmonary disease and respiratory failure
Some patients with stable COPD (those with a less intense ventilatory response to
CO2) have a high Pa,CO2. Compared with normocapnic COPD patients, hypercapnic
subjects have low Pa,O2, higher [Hb] (secondary polycythaemia) and lower resting V9E
(table 4), with shallower and more rapid breathing [6]. They are often oedematous, and
Table 2. – The common causes of respiratory failure
Primary lung failure
Adult respiratory distress syndrome
Neuromuscular
Anterior horn cell disease (e.g. poliomyelitis), phrenic nerve paresis, diaphragm myopathy
CNS failure
Brain stem (respiratory centre) depression or pathology
Multifactorial
Chronic obstructive pulmonary disease
Table 3. – Different types of respiratory failure
Cause
Pa,CO2
CO2 sensitivity
ARDS
FI,O2 need %
NIPPV response
Variable
Normal
w60%
[IPPV z PEEP]
Neuromuscular
q
Q
21–24%
z
Brain stem
q
Q
21–24%
z
COPD exacerbation
q
Q
24–35%
z
ARDS: adult respiratory distress syndrome; COPD: chronic obstructive pulmonary disease; Pa,CO2: arterial
carbon dioxide partial pressure; FI,O2: inspired oxygen fraction; NIPPV: nasal intermittent positive pressure
ventilation; IPPV: intermittent positive pressure ventilation; PEEP: positive end-expiratory pressure.
110
GAS EXCHANGE PRINCIPLES
Table 4. – Blood gases and ventilatory parameters at rest in different types of chronic
obstructive pulmonary disease (COPD) in stable state
Pa,O2
kPa
mmHg
Pa,CO2
kPa
mmHg
V9E L?min-1
f min-1
VT L
Normal
COPD "pink puffer"
COPD "blue bloater"
11–14
82–105
9.2
69
8.0
60
5.0–5.6
38–42
8.5#
14
0.62
5.3
40
9.9#
18
0.55
7.3
55
7.8#
21
0.37
Pa,O2: arterial oxygen partial pressure; Pa,CO2: arterial carbon dioxide partial pressure;
V9E: minute ventilation; f : respiratory frequency; VT: tidal volume; #: V9E would be 40%
lower if measurements were to be made without a mouthpiece and noseclip. Adapted
from GORINI et al. [6].
have been referred to as "blue bloaters". The hypercapnia of COPD patients is attributed
to alveolar hypoventilation. In this case, the hypoventilation is "functional" rather than
"actual", meaning that a large proportion of the VT is ineffective, going to units with a
high V9A/Q9 ratio, and acting as alveolar dead space. COPD is associated with severe
V9A/Q9 mismatch [7], with about one third of pulmonary blood flow going to alveolar
units with very low alveolar ventilation and contributing little to CO2 excretion (wasted
perfusion); thus, y33% of pulmonary blood flow receives 10% of total alveolar
ventilation, and the resulting low V9A/Q9 and its uneven distribution is responsible for the
hypoxaemia and hypercapnia. In fact, total V9E is normal (table 4).
In an exacerbation of COPD, Pa,CO2 rises and hypercapnia worsens, especially if
uncontrolled inspired oxygen (FI,O2 w28%) is prescribed [8]. The rise in Pa,CO2 due to the
respiratory infection alone is caused by increasing V9A/Q9 mismatch (bronchiolar
kPa mmHg
12 90
DPa,CO2
85
80
kPa
10 75 192.5mmHg
70
65
8 60
2.5 kPa
55 19 mmHg
l
(C) Unstable + O2
therapy (uncontrolled)
Pa,CO2
True hypoventilation
50
6 45
40
35
l
(B) Unstable
Worsening V´A/Q´
mismatch
DV´A
V´CO2 270 mL
0
1
l
-0.65
(A) Stable
-1.20
2
3
4
5
“Effective” alveolar ventilation L·min-1
6
Fig. 2. – Hyperbolic relationship between arterial carbon dioxide partial pressure (Pa,CO2) and "effective" alveolar
ventilation (y [VT (1 – VD/VT)] 6 f ) at constant CO2 production (V9CO2) with data from AUBIER et al. [8] in
chronic obstructive pulmonary disease (COPD) patients superimposed. As Pa,CO2 rises in acute (B) on chronic
(A) respiratory failure, so its sensitivity to a fall in alveolar ventilation increases, and small falls in tidal volume
produced by removal of hypoxic ventilatory drive (see (C)) exacerbate hypercapnia. V9A/Q9: alveolar ventilation–
perfusion; VT: tidal volume; VD: dead space. See table 5 for details.
111
J.M.B. HUGHES
inflammation and obstruction, alveolar consolidation and collapse) and a decrease in
"effective" alveolar ventilation (fig. 2, ARB). Stradling [9] has shown that the
subsequent rise in Pa,CO2 with uncontrolled O2 therapy is due to removal of the
hypoxaemic stimulus to central respiratory drive, leading to "true" alveolar hypoventilation (fig. 2, BRC). Figure 2 shows that, at a constant CO2 output, the relationship
between alveolar ventilation (for uniformly perfused and ventilated units) and Pa,CO2 is
hyperbolic, so that if Pa,CO2 is already raised (fig. 2, point B), a small fall in V9A
(0.65 L?min-1, in this example), caused by a fall in VT from 341 to 323 mL is sufficient to
raise Pa,CO2 significantly.
Oxygen carriage in blood
The ODC (fig. 1) plays an important role in gas exchange, especially in oxygen delivery
to the tissues. At a normal [Hb], w98% of O2 is Hb-bound; thus, it is convenient to
substitute for the oxygen content of arterial blood the per cent binding (saturation, S) of
Hb with O2, the Sa,O2. This presupposes that [Hb] is normal. A severely anaemic patient
may be very breathless on exercise in spite of a normal Pa,O2 and Sa,O2! In a normal lung,
the fraction of arterial oxygen carried in plasma rises from 2.5% at Pa,O2 13.3 kPa to 8%
when plasma Pa,O2 is raised to 80 kPa with 100% oxygen breathing; this reservoir of O2
(33% of requirements at rest) is a useful resource when Hb is compromised, as in carbon
monoxide (CO) poisoning.
The PO2 at half maximum blood oxygen concentration (P50)
The P50 defines the position of the ODC with respect to the PO2 axis. The normal
value for P50 is 3.5–3.7 kPa. A shift to the right (see fig. 1) unloads oxygen to the tissues
in the sense of a lower oxygen content for the same PO2; this is advantageous during strenuous exercise, and the right shift is facilitated (see fig. 1) by exercise-related
increases in tissue PCO2, hydrogen ion and temperature. 2,3-diphosphoglycerate
(2,3-DPG) is a metabolic intermediate in the glycolytic pathway; its concentration in
red cells increases in anaemia and hypoxaemia and promotes unloading of O2 to the
tissues. The shift to the left promotes oxygen loading from lung to blood (higher oxygen
content for the same PO2), and is generally considered disadvantageous; but in
circumstances where arterial hypoxaemia is very severe (and associated with
polycythaemia), such as in the foetus and at extreme altitudes, the overall effect on
tissue oxygen delivery is beneficial (see later).
Heterogeneity of ventilation and perfusion
Under nearly all circumstances (but see section Alveolar–capillary diffusion below),
alveolar PO2 and PCO2 are determined by the ratio of alveolar ventilation to perfusion
(V9A/Q9); this ratio, even in normal lungs, varies considerably from one gas exchange unit
to the next. The analysis and quantification of this heterogeneity (ygas exchange
inefficiency) is based on the PO2–PCO2 diagram (fig. 3) [10, 11]. The conceptual brilliance
of this diagram is that: 1) the V9A/Q9 line encompasses all possible V9A/Q9 values
throughout the lung (given the PO2 and PCO2 composition of mixed venous blood and the
inspired gas); 2) the blood R and gas R lines define every possible value of PO2 and PCO2
in arterial blood and mixed alveolar and expired gas; and 3) the "ideal" point defines a
"gold standard" or "perfect" lung, i.e. that value of PO2 and PCO2 the lung would have had
112
GAS EXCHANGE PRINCIPLES
kPa mmHg
6.7 50
V´A/Q´= 0
l
a
Low V´A/Q´
“shunt”
Mixed
venous ( v )
Ideal (A) point (V´A/Q´=0.86)
l
l
Pa
PA
PCO2
Blood R=0.8
E
3.3 25
0 0
l
Pe,CO2
“Shunt”
(CA–Ca)/(CA–C v )
for O2
l
High V´A/Q´
“dead space”
(A–a) PO2
“Dead space”
(Pa–Pe)/Pa
for CO2
mmHg 40
kPa 5.3
V´A/Q´ line
Gas R=0.8
Pl
70
9.3
100
13.3
130
17.3
l
V´A/Q´= ¥
160
21.3
PO2
Fig. 3. – Carbon dioxide tension (PCO2) – oxygen tension (PO2) diagram showing arterial (Pa) blood R lines and
expired gas R lines intersecting at the "ideal" point; deviations from the "ideal" point are caused by low and
high alveolar ventilation–perfusion (V9A/Q9) values in individual lung units as defined by the V9A/Q9 line. See
text for more explanation.
in the absence of V9A/Q9 heterogeneity. Another key concept is that, in the steady state,
blood and gas take up oxygen and excrete carbon dioxide in a ratio called the respiratory
exchange ratio (R), (i.e. V9CO2/V9O2), which is determined for the lung by body
metabolism, where it is called the respiratory quotient (RQ); fundamental to this idea is
that metabolism of the lung itself makes a negligible contribution to the overall gas R.
Thus, the mixed blood and the mixed gas are constrained to lines which, in relation to
mixed venous blood and inspired gas, have a fixed exchange ratio (R) (y0.8 at rest). The
blood R and gas R lines can only meet at a point (the "ideal alveolar" (A) point) where all
V9A/Q9 ratios have the same value (no heterogeneity); under resting conditions this value
is y0.86.
The composition of arterial blood is "weighted" by contributions from low V9A/Q9 units
(by definition, high V9A/Q9 units have little blood flow), and mixed alveolar (Ā9) and mixed
expired (Ē) points are similarly weighted by high V9A/Q9 units. Thus, increasing V9A/Q9
heterogeneity drives arterial pressure (Pa) and alveolar pressure (PA) (and mixed expiratory
pressure (PĒ)) values in different directions down the blood and gas R lines, and the A–a
PO2 difference becomes an index of gas exchange inefficiency. It is not possible to sample
mixed alveolar gas (Ā) because there is always contamination from the anatomic dead
space gas, so the ideal point (A) is used as the yardstick. This means that the A–a PO2
difference is weighted towards the inefficiency caused by low V9A/Q9 units.
The three compartment model of pulmonary gas exchange
This model (fig. 4) is an extension of the PO2–PCO2 diagram, using the concept of the
"ideal" point in relation to the arterial and mixed expired compositions. It is an "as if"
113
J.M.B. HUGHES
V´A
Ideal compartment
(normal gas exchange)
Alveolar dead space
(no Q´)
PCO2 ~ PI,CO2
PCO2 ~ Pa,CO2
Cc´O2
Q´ (capillary)
Q´ total
Q´s (shunt)
CV,O2
Venous admixture (no V´A)
Ca,O2
Fig. 4. – An "as if" model of lung gas exchange in which alveolar ventilation (V9A) is distributed to two
compartments (one unperfused [y alveolar dead space] and the other equally perfused and ventilated).
Pulmonary blood flow is similarly distributed to two compartments, one of which is unventilated (y shunt or
"venous admixture"). PCO2: partial pressure of carbon dioxide; Pa,CO2: arterial carbon dioxide partial pressure;
PI,CO2: inspiratory carbon dioxide partial pressure; Q (capillary): pulmonary capillary blood flow; Q9 total:
capillary and noncapillary pulmonary blood flow; Cc9O2: end capillary oxygen content; Ca,O2: arterial oxygen
content; CV̄,O2: mixed venous oxygen content.
situation. The lung behaves as if a part was "perfect" (uniformly ventilated and perfused),
defined by the "ideal" point, as if another part was perfused but not ventilated at all
(called the "physiological shunt"), and as if a third part was ventilated but not perfused
(called the "physiological dead space"). For convenience, the "ideal" point is defined in
terms of the arterial PCO2 (in figure 3, the slope of the blood R line between Pa and the
"ideal" point is fairly flat so that Pa,CO2 lies close to "ideal" PA,CO2) and an assumed value
for R of 0.8. In a simplified form for everyday use:
A{a PO2 ~½PI,O2 {Pa,CO2 =R{Pa,O2
ð2Þ
Where the first term (in brackets) is the "ideal" alveolar PO2.
Once the "ideal" point has been defined, the physiological shunt ("wasted" blood flow)
(conceptually, the distance PA – Pa in relation to PA – Pv̄) can be calculated (but in terms
of O2 contents (C) not partial pressures) as:
Q0 s=Q0 T %~½C A{C aO2 =½C A{C vO2 |100
ð3Þ
Where Q9s is the physiological shunt flow and Q9T the total pulmonary blood flow.
Dead space ("wasted" ventilation) is traditionally defined in terms of CO2 exchange,
but the principles are similar to the shunt equation; as C and P are linearly related in the
gas phase, P is retained:
V D=V T~½Pa,CO2 {PE,CO2 =Pa,CO2
ð4Þ
Where VD/VT is the physiological dead space as a proportion of the tidal volume,
114
GAS EXCHANGE PRINCIPLES
Pa,CO2 is assumed to equal the "ideal" PA,CO2 and PI,CO2 has been omitted from the
denominator.
Shunt and dead space are called "physiological" rather than "alveolar", because they
both contain an "obligatory" anatomical component, bronchial venous and Thebesian
blood flow in the case of shunt and the anatomic dead space in the case of VD/VT.
Quantitating gas exchange inefficiency for oxygen
In an earlier section, the Pa,O2 was interpreted solely in terms of the normal value for
age, BMI and posture. In equations 2, 3 and 4 (see above), gas exchange efficiency is
assessed in relation to the "ideal" or perfect lung.
The A–a PO2 gradient can be calculated from the Pa,O2 and Pa,CO2, assuming (at rest)
R=0.8. Normal values are a function of the inspired PO2 when FI,O2 i21%; for
convenience, the estimates are usually made during air breathing. The normal A–a PO2
(air) increases with age from 0.8–1.3 kPa at age 20 yrs to 3.5–4.0 kPa at age 70 yrs [12].
For the same amount of physiological shunt (Q9s/Q9T) and FI,O2, the A–a PO2 declines as
Pa,CO2 rises (and PA,O2 falls). In spite of these limitations, the A–a PO2 has been used
extensively, and for minor fluctuations in Pa,CO2 (5.33¡1.0 kPa), gives a better assessment
of gas exchange efficiency than Pa,O2 alone. In the intensive care setting, A–a PO2 is very
sensitive to FI,O2, and may increase seven-fold from 5.4 to 38 kPa, for the same Q9s/Q9T
(20%), just with an increase in FI,O2 from 21 to 60%. An empirical index, the Pa,O2/FI,O2
ratio, reduces these fluctuations, but does not abolish them.
Q9s/Q9T requires arterial O2 contents (or saturations) to be calculated from Pa,O2 and
"ideal" PA,O2 values, and an estimate made of mixed venous (pulmonary arterial) O2
content or saturation (unless right heart catheterisation has been performed). In normal
subjects at rest, an a–v difference of 50 mL?L-1 (or DSa,O2 25%) may be assumed, but in
patients with pulmonary hypertension or heart failure that assumption may be wrong.
There is probably more support for the use of the dead space – tidal volume ratio (VD/
VT); it is independent of FI,O2, and relatively independent of partial pressure of carbon
dioxide in mixed venous blood (Pv̄,CO2). The assumption that Pa,CO2=PA,CO2 will be in
error if there is a substantial Q9s/Q9T but as the a–v PCO2 difference at rest isv1.0 kPa, the
error will not be large. VD/VT is biased towards the detection of units with abnormally
high V9A/Q9. VD/VT may help in the interpretation of data, as shown in table 5, based on
data in figure 2. Normally, VD/VT at rest is v0.4. These patients with stable COPD had
hypoxaemia (Pa,O2 v8 kPa) and hypercapnia, a normal minute ventilation, but a raised
VD/VT indicating a degree of V9A/Q9 mismatch. In an exacerbation of disease, leading to
worsening respiratory failure, Pa,CO2 rose by 2.5 kPa accompanied by a rise in VD/VT and
a fall in "effective" tidal volume (VT actually rose). The fall in VT (effective) [0.11 L] was
Table 5. – Respiratory failure in chronic obstructive pulmonary disease (COPD) exacerbations; effects of
uncontrolled oxygen therapy
COPD: stable
COPD: failure
COPD: failure z O2
Resp rate
min-1
VT
L
V9E
L?min-1
VD/VT
(physiol. dead space)
VT
(effective)# L
Alveolar
ventilation}
L?min-1
Pa,CO2
[mmHg]
kPa
20
31
32
0.41
0.34
0.32
8.2
10.5
10.3
0.5
0.77
0.82
0.21
0.094
0.073
4.11
2.91
2.36
6.1 [46]
8.7 [65]
11.2 [84]
VT: tidal volume; V9E: minute ventilation; VD: dead space; Pa,CO2: arterial carbon dioxide partial pressure; #: VT
(effective)=[VT. (1 – VD/VT)]; }: alveolar ventilation=VT (effective)6resp rate. Adapted from AUBIER et al. [8] and
STRADLING et al. [9].
115
J.M.B. HUGHES
Table 6. – The causes of arterial hypoxaemia
Altitude
Low PI,O2
Hypoventilation
V9E inadequate for V9O2; Pa,CO2 always raised
Diffusion limitation
DL,O2 inadequate for V9O2; Pa,O2 falls zz on exercise
V9A/Q9 mismatch
Pa,O2 w73.3 kPa on 100% oxygen
Anatomic R–L shunt
Pa,O2 v73.3 kPa on 100% oxygen
PI,O2: partial pressure of inspired oxygen; V9E: minute ventilation; V9O2: oxygen production;
Pa,CO2: arterial carbon dioxide partial pressure; Pa,O2: arterial oxygen partial pressure; DL,O2:
oxygen diffusing capacity of the lung; V9A/Q9: alveolar ventilation–perfusion ratio; R: right; L:
Left.
much greater than the fall in VT (actual) [0.07 L], indicating severely worsening V9A/Q9
mismatch. Finally (table 5, bottom row), with uncontrolled O2 therapy, there was
another substantial rise in Pa,CO2 but very little change in VD/VT; DVT (effective) was the
same as DVT (actual), indicating that true hypoventilation was the reason for the rise of
Pa,CO2 on oxygen [6].
Causes of hypoxaemia (low arterial oxygen partial pressure) II
The cause of the hypoxaemia (table 6) is usually obvious from the clinical diagnosis.
Hypoventilation and V9A/Q9 mismatch have been discussed already. In V9A/Q9 mismatch,
Pa,O2 will only exceed 80 kPa with 15 min 100% O2 breathing if all parts of the lung are
aerated (as in COPD), with oxygen diffusing to obstructed alveoli through collateral
pathways. In ARDS or pulmonary oedema, waterlogged gas exchange units will be
unable to take up oxygen. Diffusion limitation, also called "alveolar–capillary block",
occurs on exercise at high altitude, and on exercise in patients with interstitial pulmonary
fibrosis (cryptogenic fibrosing alveolitis) whose transfer factor of the lung for carbon
monoxide (TL,CO; at rest) is v50% predicted. Intrapulmonary anatomic right to left
shunts (intracardiac R–L shunts behave similarly as regards Pa,O2) are unusual. The most
frequent causes are pulmonary arteriovenous malformations (PAVMs), associated with
hereditary haemorrhagic telangiectasia [13], and the hepatopulmonary syndrome (HPS),
associated with liver disease and portal hypertension [14]. PAVMs can be demonstrated
with high resolution computed tomography scans or pulmonary angiography. The shunt
channels in HPS are too small to be demonstrated by these techniques; contrast echobubble or albumin macroaggregate radionuclide (99mTc-MAA) lung scans will show
contrast material passing through the lung to reach the left side of the heart, or (in the
case of 99mTc-MAA) the kidneys or brain.
HPS is an interesting condition physiologically because it has features of diffusion
limitation as well as those of a R–L intrapulmonary shunt. Many of the capillaries in the
alveolar septa are remodelled (cause unknown) with diameters as large as 100–200 mM
(normal 7–15 mM). The TL,CO is very reduced (decreased capillary surface area and
increased intracapillary diffusion distances) and Pa,O2 is low and has a variable response
to breathing 100% oxygen. In HPS, a poor response to 100% O2 (Pa,O2 vv80 kPa)
suggests very widened capillaries that act as an intrapulmonary anatomic R–L shunt. On
the other hand, a good response to 100% O2 (Pa,O2 w73.3 kPa) in some HPS patients
suggests smaller channels in which diffusion equilibration can be established when the
[PA,O2–PV̄,O2] gradient is raised. With 100% O2 breathing, and arterial sampling for PO2,
the R–L shunt can be quantitated, as Qs/QT % using equation (3); this gives a
"physiological" estimate. The R–L shunt can also be measured (as Qs/QT %)
116
GAS EXCHANGE PRINCIPLES
"anatomically" using a 99mTc-MAA lung-kidney scan technique, and in large-channel R–
L shunts (as in PAVMs) these physiological and anatomic estimates are in agreement. In
HPS, the oxygen shunt (low Pa,O2) and the 99mTc-MAA shunt were the same breathing
air, but with 100% O2 breathing, the physiological shunt was less than the anatomic
99m
Tc-MAA shunt. This suggests an interesting scenario. Breathing air, the low Pa,O2 in
HPS behaves as an intrapulmonary R–L shunt, but conceptually (from the 100% O2 data)
it should be regarded as an extreme example of diffusion limitation [14, 15].
Noninvasive measurements of arterial oxygenation
The convenience of measuring arterial oxygen saturation (Sa,O2) with a finger or ear
lobe probe has been stressed earlier. The advantage of sampling arterial blood is that
Pa,CO2 and pH can also be measured. But, arterial sampling is invasive, particularly when
repeat measurements are required in ambulatory patients; in intensive care, arterial
cannulas will be inserted.
Arterialised capillary blood
A less invasive method of obtaining Pa,O2, Pa,CO2 and pH is to sample arterialised
capillary blood, obtained by making a small cut in the periphery of the ear lobe, after
previous warming with vasodilator cream. Blood, which must be freely flowing, is
collected as anaerobically as possible (with stringent precautions to avoid blood spillage
and skin pricks), and analysed immediately. Good technique is crucial. The sample is a
mixture of capillary and venular blood. The principle is that vasodilatation increases
local blood flow up to 10-fold; from the Fick equation, if local V9O2 does not change, the
arterio–venous content and PO2 difference will narrow sufficiently so that capillary and
venous PO2 approach Pa,O2. In normoxia, the a–v PO2 gradient is large (8 kPa) and
arterialised samples tend to underestimate the true arterial value (by 0.6 kPa). But, in
hypoxaemia, on the steep part of the ODC, with a smaller a–v PO2 difference (v4 kPa),
there is good convergence of arterialised PO2 and Pa,O2 at Pa,O2 levels v9.3 kPa [16]. The
results on exercise are similar to those at rest. The overall message is that false negatives
(falsely normal Pa,O2) are less of a problem than false positives, i.e. a misleadingly low
Pa,O2.
Transcutaneous measurements (Ptc,O2)
A Clark polarographic electrode placed on the skin measures the PO2 in subdermal
tissues. The principle is the same as when arterialised capillary blood is sampled.
Vasodilatation is achieved by heating the skin to 40–42uC, and this narrows the a–v PO2
difference. The method works best in neonates where the epidermis is very thin.
Substantial underestimates may occur in adults, even with gentle abrasion of the
epidermis. Calibration against a simultaneous arterial sample is needed. In adults,
transcutaneous oxygen tension (Ptc,O2) may be able to follow trends in Pa,O2 over time,
but spot samples are not reliable.
Measurement of Pa,CO2 with transcutaneous electrodes is well established as a reliable
monitor of long-term trends, i.e. overnight in patients with nocturnal hypoventilation.
The small arteriovenous difference for Pa,CO2 at rest is an advantage.
117
J.M.B. HUGHES
Table 7. – P50 (O2 partial pressure at half maximum O2 concentration), oxygen dissociation curve (ODC) shift
and haemoglobin (Hb) concentration for human blood in different situations
P50
Situation
Normal
Exercise
Hb Seattle
Hb Minneapolis
Foetal blood
Mt Everest summit
CO poisoning
HbCO 60%
kPa
3.5–3.7
3.8–4.2
5.2
2.3
2.6
2.6
1.33
ODC shift
mmHg
26–28
29–32
39
17
19
19
10
NIL
RIGHT
RIGHT
LEFT
LEFT
LEFT
LEFT
Pathogenesis
Acidosis, hypercapnia, hyperthermia
Hb variant
Hb variant
Hb variant (Hb–F)
Alkalosis#
HbCO q HbO2
antagonism; frequently fatal
[Hb]
% normal
100
100
60
117
133}
130
40
CO: carbon monoxide; Hb-F: foetal haemoglobin; HbO2: oxyhaemoglobin; HbCO: carboxyhaemoglobin; #:
alkalosis overcomes right shift effect of q 2:3 DPG (hypoxaemia induced); }: foetal umbilical blood as per cent
maternal uterine blood.
Oxygen affinity in special situations
Haemoglobinopathies
The importance of the position of the ODC, as defined by the P50 (normal value 3.5–
3.7 kPa), was stressed earlier. Shifts to the right in anaemia and hypoxaemia, produced
by an increase in red cell 2,3-DPG, promotes efficient oxygen unloading to tissues (larger
arteriovenous oxygen content difference (D[Ca–Cv̄]O2) for the same arteriovenous PO2
difference (D[Pa–Pv̄]O2)). In normoxia, shifts to the left (less O2 unloading) are
considered disadvantageous. Certain congenital haemoglobinopathies are associated
with large right or left P50 shifts (table 7). A right shift, such as occurs in Hb Seattle is
associated with anaemia (Hb 60% normal); even so, the normal a–v O2 content difference
at rest (45–50 mL?L-1) can be unloaded at a nearly normal Pv̄O2 (4.7–5.1 kPa); exercise
capacity is relatively unimpaired. On the other hand, haemoglobinopathies with a left
shift develop erythrocytosis to compensate for their difficulty in O2 unloading. Hb
Andrew–Minneapolis [17] has [Hb] 117% normal. Because of the increase in the O2
content of arterial blood, such patients can deliver 45–50 mL?L-1 to the tissues in the
normal range of Pv̄O2 (6.1 kPa).
Apart from haemoglobinopathies, shifts to the left (Q P50) occur in three other
situations: 1) CO poisoning; 2) at extreme altitudes, and 3) in the foetus. The P50 shift,
accompanied by polycythaemia, is beneficial in (2) and (3) but, accompanied by an
"effective" anaemia, it is disastrous in (1).
Carbon monoxide poisoning
Life is possible with a severe anaemia (Hb 5.8 g?dL-1: 40% normal), but replacement of
60% HbO2 with HbCO in acute CO poisoning would be fatal. The very high affinity of
CO for Hb (250 times w oxygen), actually caused by its very slow dissociation from Hb,
shifts the curve of the residual HbO2/deoxyHb to the left (a competitive antagonism
effect) so that P50 at HbCO 60% is very low (table 7); at a Pv̄O2 of 2.66 kPa, only
14 mL?L-1 of O2 would be unloaded, just 27% of the oxygen requirements at rest. In
acute poisoning, the situation is made worse by: 1) absence of a compensatory
erythrocytosis; and 2) a normal Pa,O2, and thus no ventilatory or cardiac stimulus to
tissue anoxia from the carotid body. At low levels of HbCO%, syncope is common when
mild exercise is taken because the increased oxygen demand cannot be met due to the P50
shift and the anaemia effect of replacement of HbO2 with HbCO. In severe cases, tissue
118
GAS EXCHANGE PRINCIPLES
anoxia causes loss of consciousness and ischaemic damage to the brain and heart.
Hyperbaric oxygen is an effective therapy if administered in time. At 3.0 ATM, about
50 mL?L-1 of O2 is dissolved in plasma, which is sufficient to meet O2 demand at rest [18].
The rate of dissociation of HbCO can be increased from a half time of 5 h on air to
90 min on O2 at 1.0 ATM or to 23 min at 3.0 ATM [19]. Time is of the essence, but 100%
oxygen administered in an ambulance will treat effectively those with mild CO
intoxication.
Extreme altitude
Ascending to the summit of Mt Everest (8,848 m) is an increasingly popular challenge.
Since the first ascent in 1953 by Hilary and Tensing, about 1,400 people have reached the
summit (another 180 have died on the mountain). The ascent is generally done breathing
oxygen, but a successful ascent has been made, by Messner and Habeler in 1978,
breathing air. Pulmonary gas exchange at these extreme altitudes has been studied both
in the field (American Medical Research Expedition to Everest (AMREE) 1981) and in
hypobaric chamber simulations (Operation Everest II, 1985); some measurements (PI,O2,
PA,O2, PA,CO2, heart rate) have been made standing on the summit itself [20].
The most remarkable feature of gas exchange under these extreme conditions (at the
limit of the ability of humans to cope with hypoxia [PI,O2 on the summit=5.73 kPa] is the
ability of the body to defend the arterial oxygen tension (P) and content/saturation (C,
S). The Pa,O2 on the summit was 4.7 kPa, Sa,O2 was 71% and Ca,O2 was 182 mL?L-1 (91%
of normal at sea level) [21]. The very small [PI–Pa] difference for O2 was produced by
extreme hyperventilation (at least five times normal) at rest, lowering arterial PCO2 to
1.0 kPa and raising pH to 7.7 (normal 7.4). The respiratory alkalosis shifted the
estimated P50 to 2.59 kPa and raised Sa,O2 from 40% (at normal P50) to 71% [22].
Secondary polycythaemia raised Ca,O2 from 142 mL?L-1 (at a normal Hb) to 90% of the
normal sea level value. At lower altitudes (6,300 m), resting hyperventilation was less,
Pa,CO2 was 2.45 kPa, respiratory alkalosis was mild (pH 7.47) and in vivo P50 was normal
at 3.7 kPa (the alkalosis effect being offset by a hypoxia-induced 2,3-DPG increase) [22].
The measurements of gas exchange on exercise will be considered in a later section
(section Alveolar–capillary diffusion).
Gas exchange in the foetus
The P50 of foetal Hb, and the Pa,O2 in the foetal aorta, are similar to the values on
the summit of Everest (see previous paragraph), supporting Joseph Barcroft’s 1933
description of the foetal environment as "Mt Everest in utero". Foetal [Hb] is also higher
than postnatally. There is no alkalosis (Pa,CO2 5.7 kPa [43 mmHg]), so the P50 shift is
driven entirely by the structure of foetal Hb. For an a–v PO2 difference of 45 mL?L-1,
foetal venous PO2 is low, y2.7 kPa [20 mmHg] [23]. Exercise requirements are minimal!
Diffusion
The acinus as the gas exchange unit
Experimentally, pulmonary arteries w150 mM diameter have to be blocked (with
beads) before high V9A/Q9 regions emerge [24]. 150 mM corresponds to the diameter of
the artery supplying the acinus, supporting the notion that the acinus is the effective gas
exchanging unit. There are 33–50,000 (diameter 0.06 mM) acini in the human lung [25].
119
J.M.B. HUGHES
The entry bronchiole (terminal bronchiole) branches into three generations of alveolated
respiratory bronchioles, four generations of alveolar ducts and one of alveolar sacs.
There are 250 alveolar sacs per acinus and 30–40 alveoli per sac (1,750 alveoli per acinus).
The distance from the terminal bronchiole to the alveolar sacs averages 8 mM (range 5–
13 mM); over this distance the cross-sectional area increases exponentially 64 times, like
a trumpet. During normal breathing, the acinus and its components expand and
contract, but the convective flow in and out contributes little, if anything, to the mixing
of inspired oxygen with the residual O2 present throughout the acinus at the end of the
preceding expiration; the same arguments apply, in reverse, to CO2. Alveolar ventilation,
in the sense of bringing inspired O2 molecules into contact with the alveolar epithelium,
occurs entirely by molecular diffusion, which is proportional to the physical diffusivity of
O2 in air multiplied by the surface area/distance ratio. This ratio, because of the anatomy
of the acinus, is so high (200 cm2/0.5 cm) that uniformity of alveolar PO2 has occurred by
diffusive mixing throughout the acinus by the end of inspiration [26]. Differences in
ventilation that occur because of differences in local compliance and resistance (as a
result of convective flow inequalities), cause gas concentration differences between acini,
but not within acini.
In contrast to the uniformity of acinar ventilation, acinar blood flow may be very
uneven in time and space, chiefly due to recruitment and derecruitment of pre-capillary
arterioles and alveolar septal vessels, which tend to be either "open" or "shut". This
intra-acinar non-uniformity is less evident in the dependent zones and on exercise.
Nevertheless, the uniformity of end-inspiratory PA,O2 as a result of molecular gaseous
diffusion implies uniformity of end-capillary PO2 despite non-uniform blood flow within
the acinus [26]. Thus, acinar gas exchange is determined by mean ventilation and mean
blood flow, and the resulting mean V9A/Q9 ratio. The acinus may not behave as the
ultimate gas exchange unit in disease when its architecture has been distorted, individual
alveoli flooded or alveolar-capillary membranes thickened.
Alveolar–capillary diffusion
Oxygen is transferred from gas to blood, from the alveolar epithelial surface to the Hb
molecule in the pulmonary capillary erythrocytes, according to the relationship [27]:
V 0 O2 ~DL ½PA{PcO2
ð5Þ
Where DL is the oxygen diffusing capacity (DL,O2), Pc̄ is the mean capillary PO2 and
[PA – Pc̄] is the effective (mean) driving pressure. DL,O2 is a conductance with units of
mmol?min-1?kPa-1. V9O2 [lung] must match V9O2 [body tissues]. Thus, a low exercise
DL,O2, due to interstitial lung disease (ILD), will limit V9O2,max unless the gradient [PA
– Pc̄] can be increased proportionately by increasing PA,O2 (by hyperventilation) or by
lowering Pc̄,O2 by a decrease of Pv̄,O2 on exercise. While Pc̄,O2 is one of the determinants
of V9O2,max, it is the end-capillary PO2 (Pc9,O2) which, in a uniform lung, influences the
Pa,O2. In an ideal lung (or gas exchange unit), there is diffusion equilibrium, i.e.
Pc9,O2=Pa,O2, before blood has left the alveolar region. The end gradient [PA – Pc9] for
oxygen, the existence of which implies "diffusion limitation", is a function of the
diffusion–perfusion conductance ratio:
½PA{Pc0 =½PA{Pv~e
{DL =Q0 b
ð6Þ
where [PA – Pv̄] is the initial gradient at the mixed venous entry point, DL=DL,O2, and
b for O2 is the oxygen capacitance of blood (y the slope of the ODC at any given
PO2). bO2 is high when PA,O2 and Pc9,O2 are low and the ODC slope is high. Q9b is the
perfusion conductance whose units are (if Q9 is L?min-1) mmol?min-1?kPa-1 – similar
120
PA,O2
a) 100
9.4
b) 125
l
Pc´O2
l
PA,O2
3.0
DL/Q´b
PO2 mmHg
PO2 mmHg
3.0
Normal
70
IPF (CFA) with low DL,O2
(30% pred)
40
0.4
Capillary time s
Normal
75
DL/Q´b
IPF (CFA)
REST
l
0
l
25
l
0.4
EXERCISE
l
0
0.8
Pc´O2
Diffusion limitation
GAS EXCHANGE PRINCIPLES
0.25
Capillary time s
0.5
Fig. 5. – a) Red cell oxygen partial pressure (PO2) plotted against time spent in the pulmonary capillary, starting
at the equivalent of pulmonary artery PO2 at t=0 and finishing at the end–capillary (Pc9,O2) level. The gap
[PA,O2 – Pc9,O2] represents diffusion limitation (failure to achieve complete alveolar–end capillary equilibration).
The rate of increase of red cell PO2 during capillary transit is set by the diffusion–perfusion conductance ratio
(DL/Q9b), the values being circled. b) DL/Q9b values are lower on exercise because DQ9b exceeds DDL. IPF:
idiopathic pulmonary fibrosis; CFA: cryptogenic fibrosing alveolitis; DL,O2: diffusing capacity of the lung for
oxygen; PA,O2: alveolar oxygen partial pressure.
to DL,O2; thus, DL/Q9b is the diffusion/perfusion conductance ratio. For DL/Q9b w3.0
(at rest), [PA – Pc9] isv5% of [PA – Pv̄], i.e.v0.5 kPa – almost complete equilibration.
For DL/Q9b=1.0, alveolar–capillary equilibration is only 63% complete; for a patient
with ILD on exercise, this would mean an end-gradient [PA – Pc9] of 6.3 kPa;
assuming PA,O2=13.3 kPa, Pa,O2 would be v7.0 kPa, i.e. significant hypoxaemia [28].
Figure 5 plots PA and Pc9 for oxygen and the DL/Q9b ratio (mean value for the whole
lung, ignoring regional inhomogeneity) at rest and on moderately severe exercise for a
normal subject and a patient with ILD. A small gradient opens up in the normal subject
on exercise (DQ9b (rest to exercise) wDDL). In ILD with a low DL, DL/Q9b is low at rest,
but not sufficient to cause a significant end-gradient; such hypoxaemia as exists is caused
by V9A/Q9 inequality. DL/Q9b ratio falls sharply on exercise (DL increase is small
compared to Q9 increase), causing a large [PA – Pc9] gradient ("diffusion limitation") and
exercise-induced hypoxaemia.
Diffusion limitation can occur occasionally in super-fit normal subjects breathing air,
undergoing extreme exertion when DQ9b wDDL. It occurs without exception on exercise
at altitude when PA,O2 isv8 kPa, because bO2 remains high (on the steep part of the ODC)
throughout the time course of blood capillary transit. Theoretical studies suggest that the
increase in left shift in the ODC (QP50) at altitude promotes more rapid alveolar–
capillary equilibration for any given Pb, V9O2 and DL,O2 [29], presumably by lowering bO2.
Extensive measurements of pulmonary gas exchange were made during chronic
hypobaric chamber exposure in fit subjects in Operation Everest II [30]. Diffusion
limitation was measured using the MIGET technique (see next section) by comparing the
A–a PO2 actually measured by arterial sampling with that predicted from the V9A/Q9
distribution measured by MIGET; when actual gradient wMIGET gradient, diffusion
limitation of gas exchange is inferred. Diffusion limitation was detected at sea level at
V9O2 w3.0 L?min-1, and at progressively lower V9O2s as Pb and PI,O2 decreased. On the
"summit", V9O2,max was 1.0 L?min-1 (27% of sea level value), Pa,O2 fell from 4.1 kPa to
121
J.M.B. HUGHES
3.7 kPa (rest to exercise), and A–aPO2 increased from 0.2 kPa to 0.96 kPa due to
diffusion limitation [30].
Interestingly, patients with Hb Andrew–Minneapolis with a left shift (P50 2.3 kPa) had
(at sea level) a lower V9O2,max compared with controls, but a higher V9O2,max than those at
moderate altitude (PI,O2 13.3 kPa). The authors, somewhat fancifully, termed them
"Human Llamas" [17].
Inert gas transport and the MIGET technique
The multiple inert gas elimination technique (MIGET), pioneered by Wagner et al.
[31] measures the distribution of V9A/Q9 in an "as if" 50 compartment model of the
lung; there are 48 compartments with discrete V9A/Q9 values from 0.01 to 10 plus shunt
(V9A/Q9=0) and dead space (V9A/Q9=‘) compartments. MIGET is a considerable advance
on the three compartment model (fig. 4) of Riley and Cournand [10], but it is
technologically complex and suitable only for research studies. Six inert (nonreactive
with Hb) gases with a wide range of solubilities (l), (l is similar to the capacitance
coefficient, b, except for the units (ATM-1, not kPa-1 or mmHg-1)) are dissolved and
infused intravenously for 30 min, after which mixed venous, arterial and mixed expired
blood and gas samples are taken and analysed by gas chromatography for the arterial
retention (Pa/Pv̄) and alveolar excretion (PA/Pv̄) ratio of each gas (fig. 6). The key
relationship is:
Pa=Pv~l=½lzV 0 A=Q0 ð7Þ
a) 1.0
b) 1.0
l
3 10
0.5
l
Ventilation
l
h
0.8
l
0.4
l
l
0.01 0.1
c) 1.2
a
l
0
l
l
L·min-1
0.5 V´A/Q´ 0.1 0.3 0.85
PA/Pv and Pa/Pv
Pa/Pv
l
Acetone
Ether
Halothane
Cyclopropane
Ethane
SF6
Acetone
Ether
Halothane
Cyclopropane
Ethane
SF6
Figure 6 shows that 50% retention (Pa/Pv̄=0.5) occurs with a V9A/Q9 ratio v0.1 for
0
l
A
l
1
10 100
0.01 0.1
1
Capacitance coefficient (b) mL·mL-1·Atm-1
10
100
0
Blood
flow
Shunt
l
Dead
space
l
0
0.1
1
10
Ventilation–perfusion ratio
Fig. 6. – a) Theory: each tracer gas has a unique arterial retention (Pa/Pv̄) for a given V9A/Q9 ratio. b) Retention
and excretion: values of arterial (a) retention [Pa/Pv̄] and alveolar (A) excretion [PA/Pv̄] for all six tracer gases in
a lung with moderate V9A/Q9 dispersion compared to a theoretical "ideal" lung (h) with no V9A/Q9 dispersion.
c) Analysis: presentation: plot of ventilation and blood flow (smoothed) for 48 notional compartments (plus one
each for shunt and dead space) against V9A/Q9 as a best fit to explain the data in figure b) on the basis of
theory in a).
122
GAS EXCHANGE PRINCIPLES
low solubility (l) gases, but with a V9A/Q9 ratio w10 for high solubility gases. It
follows that SF6, the gas with the lowest solubility, only has a positive Pa/Pv̄ value
from low V9A/Q9 units and shunt, for which it is the marker of choice. Conversely, the
highest solubility gas, acetone, is only retained in arterial blood from units with V9A/
Q9 w1.0, so it is a marker of high V9A/Q9 units and alveolar dead space. In figure 6b,
the overall lung retention of each gas in arterial blood (a) and alveolar gas (A) is
plotted in relation to an ideal lung (h) uniformly perfused and ventilated. The shape
of the arterial (a) and alveolar (A) lines, and the (a–h) and (h–A) pattern for the array
of inert gases (analogous to the A–a PO2) is unique for a particular V9A/Q9
distribution, which can be analysed and plotted as shown in figure 6c. The left-hand
end of the blood flow versus V9A/Q9 plot reflects poorly ventilated units, not poorly
perfused units, while the right-hand end of the ventilation versus V9A/Q9 plot
highlights units with poor perfusion. Much information about V9A/Q9 distributions
in different respiratory conditions has been obtained with the MIGET technique; an
excellent review is available [32] and West’s little book [33] is an invaluable teaching
aid.
Conclusions
Pa,O2 and Pa,CO2 are determined by several factors, principally by the properties of: 1)
blood; 2) the lung; and 3) systems controlling minute ventilation and cardiac output. The
S–shaped oxygen dissociation curve (ODC) (fig. 1) is responsible for much of the
complexity of oxygen uptake from lung to blood, its shape determining the form of
the V9A/Q9 lines and blood R in the PO2–PCO2 diagram (fig. 3). The P50 for oxygen is an
important determinant of tissue oxygen delivery (table 7). In an ideal lung, all gas
exchange units would have an optimum ratio of ventilation to blood flow (V9A/Q9)
(y0.86); heterogeneity of the ratio, due to uneven distributions of V9A and Q9, causes V9A/
Q9 mismatch and is the chief cause of arterial hypoxaemia (low Pa,O2). Respiratory failure
may occur as a result of overwhelming intrapulmonary shunt (e.g. ARDS), V9A/Q9
mismatch (e.g. COPD) or alveolar hypoventilation (extrapulmonary causes). Diffusion
limitation to gas exchange is a cause of arterial hypoxaemia in special circumstances: 1)
when DL,CO (yDL,O2) is low and cardiac output (Q9) is high (patients with lung fibrosis
exercising); and 2) in normal subjects exercising at extreme altitude.
Summary
1. The arterial oxygen tension (Pa,O2) in normal subjects is affected by several factors,
principally age, altitude and the inspired oxygen fraction (FI,O2). The arterial carbon
dioxide tension (Pa,CO2) is not affected by age, but is lowered by the hyperventilation
of pregnancy and by anxiety. In arterial blood 98–99% of oxygen is bound to
haemoglobin (Hb). Pulse oximetry is a simple noninvasive way of estimating the
oxygen saturation of Hb in arterial blood (Sa,O2) [normal=97.5%]. In anaemia, with
Hb concentration 50% normal, Sa,O2 and Pa,O2 will be normal, but arterial oxygen
content (Ca,O2) will be only 50%.
2. The commonest clinical cause (in 90% of cases) of a low Pa,O2 is uneven distribution
of alveolar ventilation (V9A) and perfusion (Q9), so-called V9A/Q9 mismatch. The
cause is intrapulmonary disease affecting the bronchi, alveoli and/ or pulmonary
circulation. The second cause (in 8%) is extrapulmonary (e.g. respiratory muscle
123
J.M.B. HUGHES
weakness, loss of CO2 chemosensitivity), involving insufficient total ventilation,
often with a tidal volume that is too small to clear the obligatory anatomic dead
space completely.
3. In chronic respiratory failure, the Pa,O2 and Sa,O2 are severely reduced (Pa,CO2 may be
low, normal or high). In acute respiratory failure, often associated with shallow
breathing and an extrapulmonary cause, the Pa,CO2 is usually raised as much as the
Pa,O2 is lowered. An increase in FI,O2 restores Pa,O2 to a "normal" level for air
breathing, whatever the cause of the respiratory failure. In the acute on chronic
respiratory failure of chronic obstructive pulmonary disease (COPD), an FI,O2
increase may exacerbate the shallow breathing and lead to a further rise in Pa,CO2.
4. The relationship between the oxygen content (CO2) of blood and its partial pressure
(PO2) – the oxygen dissociation curve (ODC) – is sigmoid in shape. The position of
the curve on the PO2 axis is defined by the PO2 at half maximum blood oxygen
concentration (ySO2 50%) - the P50). A left shift (low P50) promotes oxygen loading
in the lung, and a right shift increases oxygen unloading to the tissues. Both may be
advantageous in the right circumstances – the left shift in the foetus, and at extreme
altitude (though the left shift in carbon monoxide poisoning may be fatal), and the
right shift in strenuous exercise.
5. The normal range for Pa,O2 is quite wide. The "efficiency" of pulmonary gas exchange
is often assessed, on a quantitative basis, in terms of a physiological dead space/tidal
volume ratio (VD/VT), reflecting abnormally high V9A/Q9 ratios, physiological shunt
(Q9s/Q9T) or alveolar–arterial oxygen tension gradients (A–a PO2), reflecting the low
V9A/Q9 units.
6. Apart from V9A/Q9 mismatch and hypoventilation, a low Pa,O2 can be caused by
diffusion limitation or an anatomic shunt (either intrapulmonary or intracardiac).
The hepatopulmonary syndrome (HPS), with microvascular dilatations, is an
example of a low Pa,O2, which could be due to either or both of these causes,
depending on one’s point of view.
7. The passage of oxygen from terminal bronchioles to red cells is principally by
molecular diffusion, the final step being chemical combination with intra-red cell
Hb. The process is super-efficient, and only breaks down clinically when the surface
area for exchange is reduced by alveolar destruction (a low oxygen diffusing
capacity (DL,O2)) and pulmonary blood flow (ycardiac output) is high (e.g. on
exercise), giving a low DL/Q9 ratio.
8. The multiple inert gas elimination technique (MIGET) is a research tool for
measuring V9A/Q9 distribution in a 50 compartment model of the lung, which gives
insights into the pathogenesis of intrapulmonary disease.
Keywords: Diffusion limitation, hypoxaemia, oxygen and carbon dioxide tension in
arterial blood, partial pressure at half maximum blood concentration, respiratory
failure, ventilation–perfusion mismatch.
References
1.
2.
3.
Cerveri I, Zoia MC, Spagnolatti L, Berrayah L, Grassi M, Tinelli T. Reference values of arterial
oxygen tension in the middle-aged and elderly. Am J Resp Crit Care Med 1995; 152: 934–941.
Templeton A, Kelman GR. Maternal blood gases (PAO2–PaO2), physiological shunt and VD/VT in
normal pregnant women. Br J Anaesth 1976; 48: 1001–1004.
Rea HH, Withy SJ, Seelye ER, Harris EA. The effects of posture in venous admixture and
respiratory dead space in health. Am Rev Resp Dis 1977; 115: 571–580.
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Powers SK, Dodd S, Freeman J, Ayers GD, Samson H, McNight T. Accuracy of pulse oximetry to
estimate HbO2 fraction of total Hb during exercise. J Appl Physiol 1989; 67: 300–304.
Gorini M, Spinelli A, Ginanni R, Duranti R, Gigliotti F, Scano G. Neural respiratory drive
and neuromuscular coupling in patients with chronic obstructive pulmonary disease. Chest 1990;
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Wagner PD, Dantzker DR, Dueck R, Clausen JL, West JB. Ventilation–perfusion inequality in
chronic obstructive pulmonary disease. J Clin Invest 1977; 59: 203–216.
Aubier M, Murciano D, Milic–Emili J, et al. Effects of the administration of O2 on ventilation and
blood gases in chronic obstructive pulmonary disease during acute respiratory failure. Am Rev
Resp Dis 1980; 122: 747–754.
Stradling JR. Hypercapnia during oxygen therapy in airways obstruction: a reappraisal. Thorax
1986; 41: 897–902.
Riley RL, Cournand A. "Ideal" alveolar air and the analysis of ventilation–perfusion relationships
in the lungs. J Appl Physiol 1949; 1: 825–847.
Rahn H. A concept of mean alveolar air and the ventilation–blood flow relationships during
pulmonary gas exchange. Am J Physiol 1949; 158: 21–30.
Harris EA, Kenyon AM, Nisbet HD, Seelye ER, Whitlock RML. The normal alveolar–arterial
oxygen tension gradient in man. Clin Sci Mol Med 1974; 46: 89–104.
Whyte MKB, Hughes JMB, Jackson JE, Peters AM, Hempleman SC, Jones HA.
Cardiopulmonary response to exercise in patients with intrapulmonary intravascular shunts.
J Appl Physiol 1993; 75: 321–328.
Whyte MKB, Hughes JMB, Peters AM, Ussov W, Patel S, Burroughs AK. Analysis of right to left
shunt in the hepatopulmonary syndrome. J Hepatol 1998; 29: 85–93.
Crawford ABH, Regnis J, Laks L, Donnelly P, Engel LA, Young IH. Pulmonary vascular
dilatation and diffusion–dependent impairment of gas exchange in liver cirrhosis. Eur Respir J
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Sauty A, Uldry C, Debetaz L-F, Leuenberger P, Fitting J-W. Differences on PO2 and PCO2
between arterial and arterialised ear lobe samples. Eur Respir J 1996; 9: 186–189.
Hebbel RP, Eaton JW, Kronenberg RS, Zanjani ED, Moore LG, Berger EM. Human llamas:
adaptation to altitude in subjects with high hemoglobin oxygen affinity. J Clin Invest 1978;
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Ilano AL, Raffin TA. Management of carbon monoxide poisoning. Chest 1990; 97: 165–169.
Thom SP. Hyberbaric oxygen therapy. J Intensive Care 1989; 4: 58–74.
West JB. High Life: a history of high-altitude physiology and medicine. American Physiological
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West JB, Hackett PH, Maret JS, et al. Pulmonary gas exchange on the summit of Mt Everest.
J Appl Physiol Respir Environ Exerc Physiol 1983; 55: 678–687.
Winslow RM, Samaja M, West JB. Red cell function at extreme altitudes on Mt Everest. J Appl
Physiol Respir Environ Exerc Physiol 1984; 56: 109–116.
Longo CD, Nystrom GA. Fetal and newborn respiratory gas exchange. In: Crystal RG, West JB,
Barnes PJ, Weibel ER, eds. The Lung: Scientific Foundations. 2nd Edn. Philadelphia, LippincottRaven Publishers, 1997; pp. 2141–2149.
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Weibel ER. Design of airways and blood vessels considered as branching trees. In: Crystal RG,
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Paiva M, Engel LA. Model analysis of intra-acinar gas exchange. Respir Physiol 1985; 62: 257–272.
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Hughes JMB, Lockwood DNA, Jones HA, Clark RJ. DLCO/Q and diffusion limitation at rest and
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Wagner PD, Saltzman HA, West JB. Mesurement of continuous distributions of ventilationperfusion ratios: theory. J Appl Physiol 1974; 37: 588–599.
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126
CHAPTER 7
Transfer factor for carbon monoxide
M. Horstman, F. Mertens, H. Stam
Erasmus Medical Centre, Erasmus University, Rotterdam, The Netherlands.
Correspondence: H. Stam, Pulmonary Function Dept, Dept Pulmonary Diseases, Erasmus Medical
Centre, Erasmus University, Dr Molewaterplein 40, 3015 GD Rotterdam, The Netherlands.
The main function of the lungs is to establish gas exchange between body tissues and the
surrounding air. O2 is taken up and CO2 is eliminated.
This process of gas exchange can be subdivided into three stages:
1. Ventilation, which is the mechanism by which the alveolar gas is intermittently
freshed with ambient air. As a result the O2 concentration in the alveolar gas
remains high, and the CO2 concentration low.
2. Alveolar-capillary diffusion, which is the passive passage of gases across the bloodgas barrier.
3. Perfusion, which involves the distribution of blood in the lungs and the removal
from the lungs by the blood circulation process.
This chapter describes the characteristics of the alveolar to capillary diffusion, the
second stage in the classification above (fig. 1).
Alveolar
air
Capillary
blood
O2
Surface area
PA,O2
Pc,O2
Thickness
Fig. 1. – Schematic illustration of the O2 transfer across the gas–blood barrier. Oxygen uptake (V9O2) is
proportional to the surface area of the membrane and the partial O2 pressure gradient and inversely
proportional to the barrier thickness. Finally, V9O2 is proportional to the solubility of O2 in water and inversely
proportional to the square root of the molecular mass of O2. PA,O2: alveolar oxygen partial pressure; Pc,O2:
capillary oxygen partial pressure.
Eur Respir Mon, 2005, 31, 127–145. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
127
M. HORSTMAN ET AL.
Physiological aspects of gas exchange
In the lung the O2 transport across the gas-blood barrier per unit of time, V9O2, is:
A:K:a :
V 0 O2 !
ðPA,O2 {Pc,O2 Þ!T L,O2 :ðPA,O2 {Pc,O2 Þ
ð1Þ
d
Where: ! = proportional to
A = surface area in m2
K = diffusion coefficient of O2 in m2?s-1
d = distance in m
a = Bunsens solubility coefficient in mmol?m-3?kPa-1
PA,O2 = alveolar oxygen partial pressure in kPa
Pc,O2 = capillary oxygen partial pressure in kPa
TL,O2 = transfer factor for O2 in mmol?s-1?kPa-1
The diffusion coefficient, K, depends on the size and the mobility of gas molecules, and
therefore on the viscosity of the medium in which diffusion occurs. According to
Graham’s law the diffusion coefficient K of any gas at a specific temperature and in a
specific medium is proportional to 1/d(molecular mass). According to Forster [1]
Graham’s law is valid for respiratory gases dissolved in water. This means that V9O2 is
proportional to the pressure difference across the alveolar-capillary membrane, with a
proportionality constant TL,O2. TL,O2 is proportional to the surface area A, and inversely
proportional to the barrier thickness d. Proper gas transfer therefore requires a large
alveolar surface area and a thin gas-blood barrier. Normally, total surface area is 50–
150 m2 and the barrier thickness is y5.10-7 m.
In estimating the TL,O2, knowledge of the PA,O2 and the mean value of Pc,O2 during the
passage of blood through the capillary bed is required. There is a nonlinear increase in
the Pc,O2 of blood during passage along the capillaries, so that the difference in O2
tension across the gas–blood barrier diminishes as a function of time. In healthy
volunteers capillary PO2 equals PA,O2 after about one third of the capillary passage time.
If pressure equilibration occurs, diffusion is no longer a limiting factor and V9O2 will only
depend on the perfusion rate. Because of the non-linear increase in Pc,O2 the calculation
of the TL,O2 cannot simply be based on the mean value of mixed venous and end-capillary
PO2. Therefore, Bohr [2] and Krogh [3] suggested studying the diffusing capacity using
carbon monoxide (CO). This gas has an affinity for Hb which is y230 times larger than
that of O2. The calculation of the CO transfer factor TL,CO is based on the assumption
that the CO tension in plasma is negligible. In that case the pressure difference across the
alveolar-capillary membrane is equal to the CO tension in the alveolar gas (PA,CO) and
the transfer of CO is independent of the pulmonary perfusion rate. The TL,O2 and TL,CO
are not numerically identical. According to Krogh [3] TL,O2=1.23 TL,CO. This value of
1.23 is based on the difference in solubility and molecular mass of O2 and CO,
respectively. However, because according to Roughton and Forster [4] part of the
diffusion resistance resides within the erythrocyte and depends on the reaction rate
between CO and haemoglobin, this value 1.23 cannot be correct.
Roughton and Forster [4] described a model in which the total diffusion resistance
1/TL,CO consists of two resistances in series: the resistance of the alveolar-capillary membrane
(1/Dm) and the reactive resistance (1/hQc[Hb]) of the blood in the alveolar capillaries.
1
1
1
~
z
ð2Þ
T L,CO Dm h:Qc:½Hb
Where: Dm = membrane conductance
Qc = effective capillary blood volume, in mL
128
TRANSFER FACTOR FOR CARBON MONOXIDE
[Hb] = haemoglobin concentration as a fraction of normal
h = constant for the rate of CO uptake by the erythrocytes per mL normal blood,
in mmol?s-1?kPa-1?mL-1
The reactive resistance concerns the chemical reaction between haemoglobin and CO
and depends on the capillary blood volume and haemoglobin concentration.
While the value 1.23 is still discussed, there is consensus that it is constant and
therefore nowadays intrapulmonary gas transfer is mostly measured using CO.
Methods to determine the diffusing capacity
Several methods have been developed to estimate TL,CO. The most commonly used
techniques are the single breath, the intrabreath and multiple breath tests.
Single breath method
The single breath method is applied as follows. The patient is breathing via a two-way
valve system (fig. 2). After a maximal expiration the subject is asked to inspire as deeply
as possible a gas mixture of y0.3% CO and 5–10% helium (He) from a bag or gas
container. Flows are measured with a flow transducer (Lilly, Fleish, etc.) and the inspired
and expired volumes are obtained by integration of the flow signal in time. After a
breath-holding time of 10 s at total lung capacity (TLC) the subject exhales, and an
alveolar gas sample is collected (fig. 3). Alveolar fractions of CO and He are usually
measured in a 750 mL gas sample after discarding the first 750 mL for washout of
airways and apparatus dead space.
This technique was first described by Krogh [3]. It is based on the assumption that
after inspiring a gas mixture containing CO, the alveolar CO fraction or pressure
decreases exponentially with time during breath-holding as CO diffuses into the blood. If
the alveolar CO fraction (FA,CO) is known at the beginning and end of a time interval, it
V’
FI,CO
FA,CO
FI,He
FA,He
V
Fig. 2. – Schematic representation of equipment for measuring the single breath diffusing capacity. FI,CO: CO
fraction in the inspired gas; FI,He: helium fraction in the inspired gas; FA,CO: alveolar CO fraction; FA,He:
alveolar He fraction; V9: flow; V: volume.
129
M. HORSTMAN ET AL.
TLC
750 mL
750 mL
VC
FI,CO
FI,He
RV
Breath-holding time of 10s
Fig. 3. – Spirometric representation of the single breath manoeuvre for assessing diffusing capacity. After
exhalation to residual volume (RV) level, the patient inhales a carbon monoxide (CO) and inert gas containing
mixture to total lung capacity (TLC). After 10 s breath holding the patient exhales and an alveolar gas sample
is collected. Alveolar fractions of CO and inert gas are commonly obtained from a 750 mL gas sample after
discarding the first 750 mL for washout of anatomical and apparatus dead space. VC: vital capacity; FI,CO: CO
fraction in the inspired gas; FI,He: helium fraction in the inspired gas.
is possible to calculate the exponential decay constant (kCO) of the relationship:
F A,COt ~F A,CO0 :e{kCOðt{0Þ
ð3Þ
Where: 0 = start time in s
t = end time in s
FA,COt = FA,CO at time t
FA,CO0 = FA,CO at time 0
Forster et al. [5] modified the single breath technique by adding the inert gas He to the
inspired gas mixture. They measured the He fraction both in the inspired gas, and in the
expired gas. Assuming He is insoluble in blood and tissues, they calculated alveolar
volume (VA) from the He dilution and the inspired volume (VI). In a mass balance the
total volume of He in VA is equal to the inspired volume of He:
V A:F A,He~F I,He:ðV I{V DÞ
ð4Þ
Where: VA = alveolar volume in litres BTPS (at body temperature and ambient
pressure, and saturated with water)
FI,He = He fraction in the inspired gas
FA,He = alveolar He fraction at time t
VI = inspired volume in litres BTPS
VD = total dead space in litres BTPS
Because the He analyser is sensitive to CO2, the last is absorbed prior to both He and
CO analysis. The remaining gas concentrations are usually corrected for an absorbed
amount of 5% CO2 [6].
Usually V I is equal to the inspiratory vital capacity (IVC), and the maximum alveolar
volume (VA,max) is calculated according to:
F I,He :
V A,max~
ðIVCV DÞ
ð5Þ
F A,He
Forster et al. [5] assumed that He and CO are diluted in a comparable way, which is
130
TRANSFER FACTOR FOR CARBON MONOXIDE
still generally accepted. Then, the initial fraction of CO (FA,CO0) can be approximated
from the measured inspired CO fraction and the degree to which He is diluted by residual
volume (RV), according to:
F A,He F A,CO0
~
ð6Þ
F I,He
F I,CO
Where: FI,CO = CO fraction in the inspired gas
FI,He = He fraction in the inspired gas
FA,He = alveolar He fraction (after t sec)
FA,CO0 = alveolar CO fraction at zero time
Another modification was made by Jones and Meade [7], who demonstrated that the
effective breath-holding time was not equal to the time the subjects held their breath at
TLC. The effective breath-holding time starts when 1/3 of the vital capacity is inspired
and lasts until half of the alveolar sample is collected.
In equation (3) kCO (s-1) represents:
T L,CO:ðPB{PH2 O,satÞ
kCO~
ð7Þ
KSTPD:V A,max
Where: TL,CO is in mmol?s-1?kPa-1
PB = barometric pressure and PH2O,sat = the saturated water vapour pressure at
body temperature (usually 37uC) both in kPa
VA,max = the alveolar volume at TLC level in litres BTPS
KSTPD = the conversion factor for the conversion from litres BTPS to STPD (a
volume of gas at standard temperature of 0uC and pressure of 760 mmHg
that contains no water vapour) and from L to mmol.
Equation 3 can be rewritten as:
F A,CO0
T L,CO:ðPB{PH2 O,satÞ :
~kCO:ðt{0Þ~
ln
t
KSTPD:V A,max
F A,COt
ð8Þ
Rearrangement gives:
1
KSTPD
:ln F A,CO0
T L,CO~V A,max: :
F A,COt
t ðPB{PH2 O,satÞ
ð9Þ
The exponential decay constant kCO is the primary variable; it is proportional to
TL,CO/VA. TL,CO is therefore obtained by multiplying TL,CO/VA with VA. Both TL,CO
and TL,CO/VA are used to describe the diffusion properties of the gas–blood barrier.
Ogilvie et al. [8] described the single breath method in detail with respect to dead space
wash-out volume, breath-holding time, effects of changes in intrathoracic pressure, body
position and lung volume and studied the reproducibility of the test. The modern single
breath test is based on Ogilvie’s paper and on European Respiratory Society (ERS) [9]
and American Thoracic Society (ATS) [10] guidelines. To minimise variability the ERS
and ATS give recommendations to deal with factors that affect pulmonary capillary
blood volume, CO back tension, submaximal inspired volume, prolonged inspiration or
expiration times and not optimal breath-holding conditions. Several of these sources of
error are discussed in the section entitled "Factors influencing the diffusion
measurement".
Three equations method. The conventional single breath method assumes fast
inspiration and expiration. In the case of reduced inspiratory and/or expiratory flows,
the accuracy and reproducibility of the single breath test are improved by implementing
131
M. HORSTMAN ET AL.
the three equations method (DL,COSB-3EQ) [11]. When using rapidly responding CO and
inert gas analysers, different algorithms can be used for inhalation, breath-holding (the
Krogh equation [3]) and exhalation, respectively. Such a refinement makes the single
breath test a more useful marker of disease, in particular in obstructive patients who
inhale and exhale slowly. Graham et al. [12] reported an improved precision and accuracy
of TL,CO estimates using the DL,COSB-3EQ method. The ATS Epidemiology
Standardization Project [6] recommended this technique when single breath
manoeuvres are performed with reduced flows and/or at reduced breath-holding time.
The DL,COSB-3EQ method appears to be comparable with the traditional single breath
test when inhalation and exhalation are forced and breath-holding time is y10 s.
Advantages and limitations. The single breath method is considered the "gold standard"
to determine transfer factor. The fact that breath-holding occurs at TLC level is an
advantage as well as a disadvantage of the single breath method. The advantage is that
TLC is a reproducible reference point. A disadvantage of breath-holding at TLC is that
diffusing capacity is measured at a non-physiological lung volume. Another disadvantage
is that not every patient is capable to perform the single breath procedure. Either the
patient cannot hold his breath at TLC for 10 s, or cannot deliver the required 1.5 L
exhaled volume (0.75 L for washout of dead space and 0.75 L alveolar gas sample).
Traditionally the single breath method utilises a single alveolar gas sample, which is
assumed to be representative of the entire lung. Implicitly the lung is assumed to be one
single, well-mixed compartment with one TL,CO and TL,CO/VA value, respectively.
However, CO uptake occurs in a large number of acini, each with their own relative
contribution. In obstructive patients inhaled CO will be preferentially distributed to the
better-ventilated lung areas and the single breath transfer factor will accordingly be
weighted towards these well-ventilated areas.
Predicted values single breath transfer factor. The ERS [9] reported predicted values
for Caucasians, which are dependent on age, stature and sex. The predicted values for
TL,CO were derived from studies carried out with comparable equipment and techniques,
which seemed to be compatible with the recommendations. The equations are a summary
of the mean from literature. The corresponding TL,CO/VA predicted values should be
calculated from TL,CO and TLC predicted values. The ethnic component of the transfer
factor is small and for clinical purposes unimportant [9]. When using a specific set of
predicted values, investigators should be sure the measurement conditions are comparable
with laboratory conditions (e.g. percentage of O2 in the inspiratory gas mixture). Stam
et al. [13] described predicted values for children from 6–18 yrs of age. TL,CO increases
and TL,CO/VA decreases exponentially with height. Because TLC is also exponentially
related with height, both TL,CO and TL,CO/VA are linearly related with TLC.
Intrabreath method
The intrabreath method attempts to obtain information on the distribution of
TL,CO/VA. One uses a rapid responding CO and CH4 analyser [14–16] or a mass
spectrometer for fast He analysis. After a maximal inspiration TL,CO is measured
continuously after a brief breath-holding time (1–2 s) during one single, slow and
maximal exhalation performed at a relatively constant flow. A flow restrictor and/or an
on-screen flow indicator can be used to maintain the desired flow (0.3–0.6 L?s-1 or 0.5–
1 L?s-1). Using the traditional TL,CO equation (9), TL,CO is repeatedly calculated during
the entire exhalation manoeuvre from 10% increments of exhaled volume using the CO
fractions at the beginning and end of each volume increment.
132
TRANSFER FACTOR FOR CARBON MONOXIDE
Advantages and limitations. An advantage is that this method will result in a TL,CO/VA,
which is varying during the exhalation and which might explain regional differences in
diffusion characteristics. Furthermore, a vital capacity v1.5 L is no longer a limitation,
and the breath-holding time needs to be brief only. A disadvantage is that not every
patient will be able to produce a low and constant expiratory flow. The use of a flow
restrictor to obtain a constant expiratory flow has the disadvantage of increasing intrathoracic pressure if the expiration attempt is too forced, causing a decreased effective
capillary blood volume and therefore a decreased diffusing capacity.
Multiple breath methods
In children and very ill adults, the single breath manoeuvre is difficult to perform.
Multiple breath methods have been developed to avoid the necessity of breathholding
manoeuvres and a minimum vital capacity (VC) of 1.5 L.
Steady state method. In the steady state technique (Filley et al. [17] and Bates et al. [18])
subjects breathe during a certain time from a gas container filled with a gas mixture with a
low CO concentration. Mixed expired CO is monitored until a steady state is reached. The
diffusing capacity under steady state conditions is estimated from:
V 0 CO
ð10Þ
PA,CO
Where V9CO is the CO uptake, which is calculated from inspired and expired amount
of CO, and PA,CO is the alveolar CO tension. The reproducibility is low and its
importance is limited because the results of this method depend on the minute
volume. Furthermore, the CO load of the measurement is high.
T L,CO~
Rebreathing method. Another multiple breath method is the rebreathing technique
introduced by KrÜ hoffer [19]. The subjects hyperventilate for y30 s from a bag
containing a gas mixture with a low CO and inert gas concentration. Breathing must be
performed at a large tidal volume and at a rate ofy30 breaths per minute. The gas in the
lungs is assumed to be well mixed with the gas in the rebreathing device. An inert gas is
added to measure lung volume and the total volume of the rebreathing system. TL,CO is
calculated from initial and final CO fractions, comparable to the single breath method. As
in the steady state method results of this method are also dependent on breathing pattern.
An advantage over the steady state method is the smaller CO load, because the inspiratory
CO fraction decreases during the measurement.
Usually the rebreathing technique does not entail CO2 absorption or O2 supplementation. As a consequence, the measurement time must be brief. Furthermore, the measurements are usually performed during voluntary hyperventilation to approximate one
compartment for the alveolar volume, the dead space and the volume in the rebreathing
device. Patients who are too ill to perform a single breath test will also have problems
with such a hyperventilation procedure. Therefore, Stam et al. [20] developed a
rebreathing method at normal resting ventilation in which CO2 is absorbed and O2
supplied. The system consists of a bellows in which the O2 concentration is kept between
20–22% (figs 4 and 5). The ventilation of the subject is measured with a displacement
transducer connected to the bellows. He, CO and O2 concentrations are analysed
continuously. The patient is connected to the rebreathing system at functional residual
capacity (FRC) level. During the measurement the CO fraction decreases both by
dilution with alveolar gas and by diffusion across the alveolar-capillary membrane. FRC
is estimated by monitoring the dilution of He (fig. 6). By comparing the exponential
133
M. HORSTMAN ET AL.
Displacement transducer
Sodalime
CO2 absorber
Bellows
Ventilator
He, CO and O2
analysis
O2 suppl.
Dead space
20 mL
Valve
Fig. 4. – Schematic representation of the rebreathing system to measure diffusing capacity during rest ventilation.
The disappearance of carbon monoxide (CO) from the system and dilution of helium (He) are monitored
continuously.
Fig. 5. – Measurement of the rebreathing diffusing capacity. The child breathes quietly into a closed bellows system
filled with a gas mixture containing CO and He. The picture is published with permission of the parents and child.
decay of He and CO in time, the dilution independent time constant of the CO
disappearance, kCO, can be calculated from the linear decay of the natural logarithm of
the CO fraction in time (fig. 7). This time constant is representative of the average TL,CO/
VA during resting ventilation. Prediction equations in adults are based on VA, alveolar
ventilation (V9A) and age, in children on VA, V9A and height.
Advantages and limitations. An advantage of the rebreathing technique developed by
Stam et al. [20] is that the transfer factor can be obtained in patients with very small lung
134
TRANSFER FACTOR FOR CARBON MONOXIDE
5
0.4
0.3
3
0.2
2
0.1
1
0
CO %
He % and volume L
4
0
50
100
Time s
150
200
0
Fig. 6. – Helium (He) dilution and exponential decrease in carbon monoxide (CO) concentration during a
rebreathing measurement. - - -: He concentration; ––: CO concentration; ??????: volume change (position of
bellows).
0
-0.5
In CO
-1.0
-1.5
-2.0
-2.5
-3.0
-3.5
0
50
100
Time s
150
200
Fig. 7. – Linear decay of the natural logarithm of the carbon monoxide (CO) concentration in time. The
exponential decay constant, kCO, is calculated from the slope of this relationship.
volumes while breathing normally. Because kCO depends on V9A, minute ventilation
needs to remain as constant as possible during the test.
Determining the components of the transfer factor
As Roughton and Forster [4] described in their model (equation 2), the total
diffusion resistance, 1/TL,CO, is subdivided in its components 1/Dm and 1/hQc [Hb].
Because the rate h for the reaction between CO and Hb depends on the O2 tension, TL,CO
also varies with the O2 tension. The estimation of Dm and Qc is based on single breath
TL,CO measurements at two different O2 levels and thus two different reaction rates. The
135
M. HORSTMAN ET AL.
values of 1/TL,CO are linearly related with 1/h with constants for 1/Dm and 1/Qc. 1/Dm
and 1/Qc are determined from the intercept with the ordinate and the slope of the linear
relationship, respectively (fig. 8). According to Roughton and Forster [4] h can be
calculated from the ideal alveolar oxygen tension.
Simultaneous CO and NO measurement. Borland and Higgenbottam [21] and
Guenard et al. [22] described another technique to separate the transfer factor into its
components. They measured TL,CO and transfer factor for NO (TL,NO) simultaneously,
using the single breath method. 1/TL,CO is influenced by both the membrane resistance
1/Dm and the reactive resistance 1/hQc. Since NO reacts much faster with haemoglobin,
TL,NO is much less influenced by the reaction rate with haemoglobin and is therefore a
good estimate of Dm. As a result, it is possible to calculate Dm and Qc from simultaneous
CO and NO measurements.
Advantages and limitations. An advantage of determining the components of the
transfer factor is that it is possible to differentiate whether diffusion disturbances are
caused by interstitial or capillary pathology. TL,CO can be lowered because of an
alteration in Dm, in Qc or a combination of both. The accuracy of the estimation of Dm
with the graphical method with various O2 tensions is limited, because 1/Dm is about zero.
A normal variation in TL,CO at both O2 levels might lead to an admittedly small variation
in 1/Dm, but a large variation in Dm. It is even possible to obtain a negative Dm. A
limitation of the simultaneous CO and NO method is the fast disappearance of NO.
Breath-holding time has to be reduced to ƒ5 s for the estimation of TL,NO, while for the
estimation of TL,CO a breath-holding time v5 s is too short.
Factors influencing the diffusion measurement
The Roughton and Forster [4] model clearly illustrates that the diffusing capacity
depends on: haemoglobin concentration, oxygen tension and effective capillary blood
volume. Therefore, an estimate of the diffusing capacity without knowledge of the Hb
concentration is of limited value, and a Hb correction is always required. The diffusion
indices are corrected for abnormal Hb concentrations according to the procedure
l
l
l
1/TL,CO
Slope: 1/Qc
l
l
l
Intercept: 1/Dm
1/0
Fig. 8. – Graphical determination of the components of carbon dioxide transfer factor (TL,CO), i.e. membrane
conductance (Dm) and effective capillary blood volume (Qc). 1/Dm is represented by the intercept with the
ordinate and 1/Qc by the slope of the linear relationship between 1/ TL,CO and 1/h. h: constant for the rate of
CO uptake by the erythrocytes.
136
TRANSFER FACTOR FOR CARBON MONOXIDE
described by the European Community for Coal and Steel (ECCS) [9].
azh:½Hb
T L,COðcorrÞ~T L,COðobsÞ:
ðazhÞ:½Hb
ð11Þ
Where: TL,CO (corr) = TL,CO corrected to reference Hb concentration
TL,CO (obs) = the observed TL,CO at the actual Hb concentration
a = the ratio of membrane conductance and capillary blood volume (in traditional units mL, min and mmHgy0.7 and in SI units mmol, min and kPa 230)
h = the reaction rate for the CO Hb reaction at an oxygen pressure of
110 mmHg
[Hb] = haemoglobin as a fraction of normal.
The O2 dependence of the reactive resistance implies that patients on supplementary
O2 have to be disconnected from the oxygen supply for at i10 mins prior to a diffusion
measurement [9]. Arising from the dependence of the transfer factor on the effective
capillary blood volume, the body position needs to be upright during the measurement,
because reference values were determined in that posture. The lung is more equally
perfused in the supine position, resulting in a larger effective capillary blood volume and
a larger diffusing capacity [23]. This might be due to a shift of blood from the systemic
circulation into the pulmonary circulation when changing from upright to recumbent
posture. According to Lewis et al. [24], capillaries are simply endothelial tubes that open
fully if transmural pressure exceeds a critical opening pressure. As a consequence
capillaries in the basal parts of the lungs in the sitting position will be fully open, whereas
in the apex the majority of the capillaries are closed. In the supine position gravitational
effects have less effect, resulting in a more uniform perfusion and therefore in a larger
effective capillary blood volume Qc (Bryan et al. [25] and Stam et al. [23]). These data
suggest that the lung is an overdimensioned gas exchanger with a large reserve of
capillary blood volume. An increase in effective Qc results in an increased TL,CO and
TL,CO/VA in the supine position. A lack of response to a change in body position in
various pulmonary or cardiac diseases seems to be an indication that the capillaries in the
upper lung zones are fully recruited in both positions [26, 27]. When cardiac output is
increased due to physical activity, effective capillary blood volume is also increased due
to distension and recruitment of capillaries. Therefore, a patient needs to take a rest for
y5 min before starting the diffusion test.
Furthermore, alveolar pressure should be near atmospheric during the breath-holding
time. A Valsalva manoeuvre decreases and a Muller manoevre increases capillary blood
volume and therefore TL,CO and TL,CO/VA [28].
Furthermore, the results of a diffusion measurement are influenced by: alveolar
volume at which the measurement is performed; the CO back tension in the capillary
blood; and washout time of test gases in between the various measurements.
As mentioned, the TL,CO is proportional to the surface area A of the blood–gas
barrier. A decrease in lung volume will cause a decrease in surface area A and
consequently, in TL,CO. However TL,CO/VA is higher at reduced alveolar volumes,
compared with reference values estimated at a normal TLC [9], because VA (proportional
with the radius to the third power) is decreasing faster than TL,CO (proportional with
surface area and thus with the radius to the second power). Therefore, it is important that
the inspired volume during the single breath procedure is as close as possible to the
known VC.
Because transfer factor is measured using CO and it is assumed that the capillary CO
pressure equals zero, the number of successive single breath tests is limited to a maximum
of five measurements a day. If, for some reason, this number is exceeded, corrections
should be made for CO back tension. Failing to correct for back tension, in smokers the
137
M. HORSTMAN ET AL.
transfer factor will be underestimated. Similarly, after a recent cigarette CO back tension
correction is required or the test should be postponed.
In between measurements, a minimal interval of 4 min is required to allow elimination
of test gases from the lung. During this interval the patient should remain at rest and
seated.
Comparison of single breath and rebreathing method during
rest ventilation
The absolute values of TL,CO and TL,CO/VA obtained with the various methods are not
the same. The main reason is that with the single breath method diffusion parameters are
estimated at TLC, whereas with the steady state or rebreathing methods they are
estimated at a smaller lung volume (FRCz1/2 tidal volume).
As described by Stam et al. [13, 29, 30] there is a linear relationship between single
breath TL,CO/VA and VA. Extrapolation to the lung volume range at which the various
rebreathing measurements were performed results in the shaded area in figure 9a.
Because CO disappears in the alveoli only, the VA, VD and rebreathing system
constitute separate compartments. Theoretically, these compartments can be regarded as
one compartment at infinite ventilation only. Figure 9b illustrates that above an alveolar
ventilation of 30 L?min-1 the absolute values of rebreathing TL,CO/VA are comparable
with those of the single breath TL,CO/VA at the same volume level.
Therefore, predicted values of rebreathing diffusing capacity during rest ventilation
are not only age dependent, but should also depend on alveolar volume and alveolar
ventilation.
Reporting of results and interpretation
It is important that the report includes the following data for optimal interpretation
(table 1). The diffusion measurement should be performed at least twice. At Erasmus
University in Rotterdam at least three measurements are performed (columns 1, 2 and 3)
with the average in column 4. Column 7 is the predicted value and the last column is the
standard deviation in the predicted value. In column 5 the percentage of predicted and in
column 6 the standard deviation score (SDS) or Z-score is reported. The first four rows
are concerning the volumes at which the single breath test is performed. The next four
rows give TL,CO and TL,CO/VA respectively, in which the values with the subscript c
correspond with the diffusion indices corrected to a normal Hb concentration. As
mentioned earlier, a correct interpretation of the diffusion data is only possible if the Hb
concentration is known. In row 9 TL,CO/VA is not compared with a predicted value at
predicted TLC, but with a TL,CO/VA predicted value at the actual TLC [29].
In row 4 the measured VC during the single breath manoeuvre (VCsb) is compared
with the known VC from spirometry (VCspir). If the patient exhaled and inhaled
maximally the ratio between VCsb and VCspir isy1. Ventilation distribution unequality
is evaluated based on the ratio between TLC determined with the single breath test
(TLCsb) and TLC determined with the multiple breath He washin method (TLCmb). A
TLCsb/TLCmb ratio w0.85 has been regarded as an indication for normal ventilation
distribution (Roberts et al. [31]). Conclusions about unequal ventilation are only valid
when the VCsb/VCspir ratio is y1. If the ratio VCsb/VCspir v1 and the TLCsb/TLCmb
v0.85, then unequal ventilation cannot be excluded, but TLCsb may be measured partly
too low due to submaximal inspiration. If VCsb/VCspirv1 and TLCsb/TLCmb=1, then
138
TRANSFER FACTOR FOR CARBON MONOXIDE
a) 50
TL,CO/VA mmoL·s-1·kPa-1·L-1
45
40
35
n
30
n
n
n n
25
n
n
n
nn
20
nn
15
10
5
0
VA,r
0
1
2
3
b) 50
4
VA L
5
6
7
8
TL,CO/VA mmoL·s-1·kPa-1·L-1
45
40
35
30
25
s
s
ss
s
s
s
s
20
s
15
s
s
10
5
0
0
5
10
15
20
25
30
V´A L·min-1
35
40
45
50
Fig. 9. – The single breath TL,CO/VA as function of VA (a) and the rebreathing TL,CO/VA as function of alveolar
ventilation (V9A) (b) in a healthy volunteer. VA,r is the alveolar volume range between the mean ¡2 SD
obtained from the rebreathing manoeuvres. The shaded area in a) represents the range in single breath
TL,CO/VA corresponding to the volume range of VA,r. This area corresponds with the shaded area in b).
The dashed line in b) is the linear regression line for the TL,CO/VA versus V9A relationship up to a V9A of
20 L?min-1.
Table 1. – Layout of report of the diffusion measurement
TLCsb L
TLCsb/TLCmb
VCsb L
VCsb/VCspir
TL,CO mmol?s-1?kPa-1
TL,COc mmol?s-1?kPa-1
TL,CO/VA mmol?s-1?kPa-1?L-1
TL,CO/VAc mmol?s-1?kPa-1?L-1
TL,CO/VAc RCL mmol?s-1?kPa-1?L-1
Hb mmol?L-1
1
2
3
Average
% pred
Z-score
Pred
SD
6.15
6.22
6.22
89
-1.13
6.98
0.70
4.85
4.91
4.90
106
0.46
4.63
0.56
43.02
43.02
7.18
7.18
46.96
46.96
7.75
7.75
48.84
48.84
8.06
8.06
6.19
0.69
4.89
0.98
46.32
43.90
7.68
7.28
29
27
34
32
32
-4.62
-4.72
-4.47
-4.59
-4.66
161.9
161.9
22.67
22.67
23.93
25.00
25.00
3.35
3.35
3.49
10.30
TLC: total lung capacity; sb: single breath test; mb: multiple breath test; VC: vital capacity; spir: spirometry; TL,CO:
transfer factor for carbon monoxide; c: corrected to standard Hb concentration; VA: alveolar volume; TL,CO/VAc
RCL: predicted value at actual TLC.
139
M. HORSTMAN ET AL.
Table 2. – Deviation from predicted values each measured value
Normal
Mildly decreased
Moderately decreased
Moderately severe
Severe
v measured value v
w measured value w
w measured value w
w measured value w
w measured value
-1.64 SD
-1.64 SD
-2 SD
-3 SD
-4 SD
z1.64 SD
-2 SD
-3 SD
-4 SD
VCsb is too small at the expiration side (residual volume is not reached). In that case the
single breath test is performed at TLC level and the TL,CO and TL,CO/VA results are not
influenced by this decreased VCsb.
Most pulmonologists use percentage of predicted when comparing the results with
predicted values. The diffusion indices are normally distributed and a normal range
between z1.64 SD and -1.64 sd from predicted is assumed (90% of the healthy
volunteers). A SDS- or Z-score, i.e. the deviation in sd from predicted ((measuredpredicted)/sd), is used to detect the severity of the pathology. The Z-score is related to the
chance that the index is normal. In the Erasmus University in Rotterdam it is agreed that
a deviation from predicted is judged as in table 2.
In case of a TLCsb/TLCspir w0.85 (equal ventilation distribution) and a decreased
TLCsb, TL,CO/VA, is compared with predicted values at predicted TLC (row 8), as well
as with predicted values at the actual disease limited TLC (row 9).
Clinical indications
Chronic obstructive pulmonary disease
Chronic obstructive pulmonary disease (COPD) is one of the major causes of death
worldwide. Loss of alveolar surface area and dysfunction of the alveolar membrane as in
emphysema lead to a decreased transfer factor. Measurement of the transfer factor can
be of importance in the (early) detection of COPD.
Interstitial lung disease
Thickening of the alveolar membrane and a diminished total lung capacity due to
interstitial processes may lead to a severe decline in transfer factor. The acinus is
disrupted and the diffusion pathway is lengthened. Typical diseases are extrinsic allergic
alveolitis, pulmonary vasculitis syndromes, systemic lupus erythematosus, and of course,
interstitial fibrosis.
Pulmonary bleeding disorders
Measurement of the transfer factor can be helpful in detecting intrapulmonary
bleeding in patients with disorders such as primary pulmonary haemosiderosis,
Wegener’s disease or Goodpasture’s syndrome. Because of the high affinity between
CO and Hb the TL,CO and TL,CO/VA can be increased appreciably in alveolar
haemorrhage, because CO will react with Hb without the need to pass the gas–blood
barrier. A typical feature is the gradual decrease in diffusing capacity after several
measurements, because the blood in the alveoli becomes saturated with CO.
140
TRANSFER FACTOR FOR CARBON MONOXIDE
Pre-operative screening
It is important to screen the transfer factor pre-operatively prior to any major surgery
to predict whether problems can be expected during anaesthesia or in the post-operative
phase. It is also recommended to measure the transfer factor prior to any lung surgery
(e.g. resection due to lung cancer), because resection will result in loss of surface area.
Interpretation of diffusing capacity in:
Restrictive disease
TL,CO and TL,CO/VA are usually compared with predicted values, which are
determined in healthy volunteers, who by definition have a normal TLC. Thus the
current predicted values relate to measurements made at normal TLC [9]. In patients
with a restrictive ventilatory defect (i.e. a reduced TLC) or with a larger than normal
TLC, a comparison with predicted values at predicted TLC can lead to erroneous
conclusions. A decrease in lung volume will cause a decrease in surface area A and
consequently in TL,CO. However TL,CO/VA is higher at reduced alveolar volumes,
compared with predicted values estimated at a normal TLC, because VA (proportional to
the radius to the third power) is decreasing faster than TL,CO (proportional to the surface
area and thus to the radius to the second power).
Stam et al. [29] suggested that in restrictive pulmonary disease TL,CO and TL,CO/VA
should be compared with predicted values at a lung volume equal to the patients actual
TLC. Therefore, they derived reference values for TL,CO/VA as a function of alveolar
volume. Their results were corroborated by Chinn et al. [32] and Frans et al. [33], who
found a comparable relationship between TL,CO and VA. Johnson [34] proposed a
procedure to correct predicted values of TL,CO and TL,CO/VA at predicted normal TLC
to a symptom limited TLC. However, a disadvantage of such a method is that both
predicted values of TL,CO and TL,CO/VA at predicted TLC, and the volume correction
procedure, have their own variability. The standard deviations of the calculated TL,CO
and TL,CO/VA predicted values at lower VA levels are considerably larger than at normal
TLC, and therefore conclusions concerning the severity of pathology are more difficult.
However, Hughes et al. [35] criticised a voluntary volume reduction model in normal
subjects. They stated that this model describes a restriction of extra pulmonary origin or
due to respiratory muscle weakness only. Because this model assumes uniform changes
and in interstitial pulmonary disease structural and functional changes are nonuniform,
they stated that a restriction due to interstitial lung disease is not comparable with
voluntary volume reduction in normals. However, Stam et al. [30] studied a group of
males without previous pulmonary disease before and after treatment with bleomycin for
a germ cell tumour. All of the studied subjects developed a diffusion disturbance and half
of them also developed a restriction. It was observed that the slope in TL,CO and TL,CO/
VA with change in VA was similar before and after the treatment with bleomycin. This
supports the contention that the extent of diffusion disturbance was assessed more
correctly, and appeared greater, if TL,CO and TL,CO/VA were compared with reference
values at actual TLC, rather than to values at predicted or pretreatment TLC.
Furthermore, Hughes et al. [35] stated that voluntary volume reduction is not
comparable with pneumonectomy. They described another model for loss of alveolar
units, which is based on the work of Hsia et al. [36], who described an increase in
TL,CO/VA during exercise due to a larger pulmonary blood flow, causing a more equal
perfusion by capillary recruitment. Hughes et al. [35] assumed that total pulmonary
141
M. HORSTMAN ET AL.
blood flow remains at pre-resection level, so that after resection the flow to the remaining
lung will increase about two-fold. They describe that this situation is comparable with the
dependance of TL,CO/VA on cardiac output as described by Hsia et al. [36]. Corris et al.
[37] established an empirical relationship for the increase in TL,CO/VA (post-pre
pneumonectomy) based on the percentage of flow to the resected lung pre-operatively.
These predictions are comparable with the Hughes model.
It is important to use an appropriate model for reference values depending on the
origin of the restriction. In interstitial pulmonary disease an appropriate model is not
obvious, but comparison with predicted values at predicted TLC will lead to an
overestimation of TL,CO/VA. Therefore, it is important to perform more extensive
research in this particular field.
Obstructive disease
In healthy volunteers the assumption that a small sample of air early during the
exhalation is representative of the entire lung seems to be acceptable, but in patients with
uneven ventilation and uneven distribution of TL,CO/VA the analysis of only one small gas
sample might lead to erroneous conclusions. This is because primarily the CO uptake in the
well ventilated parts of the lung will be estimated, while a different TL,CO/VA can be
expected in the poorly ventilated lung areas. An indication of this ventilation unequality is a
ratio TLCsb/TLCmb, which is v0.85 [31]. Not only unequal ventilation, but uneven
distribution of gas transfer TL,CO/VA might occur. TL,CO/VA will change during the
exhalation and here methods such as the intrabreath method come into play [14–16].
Improvement and validation of diffusion equipment
Worldwide, there are many manufacturers of equipment for measuring the transfer
factor of the lung. Diffusion data measured with the apparatus of the various
manufacturers differ significantly. Recently, Gissmeyer et al. [38] and Jensen and Crapo
[39] developed a single breath TL,CO simulator. By producing gas mixtures of CO and an
inert gas with air, this equipment creates a constant and adjustable TLC, TL,CO and
TL,CO/VA. A diffusion simulator will not only be valuable in comparing equipment, but
it will also be valuable in the regular calibrating of equipment instead of the customary
biological calibration.
Conclusion
The transfer factor of the lung has become a major lung function index and is an
important diagnostic index in COPD, interstitial pathology, etc. Several methods to
determine the transfer factor were developed in the last century. Each method has its own
advantages and limitations. The single breath method became the most generally
accepted method worldwide. Standardisation is important to diminish the variability of
the single breath method and, therefore, the ERS and ATS recommended guidelines.
However, in the case of unequal ventilation or unequal distribution of diffusion
characteristics the traditional single breath test is insufficient. One of the possible prospects,
when using fast responding gas analysers, is to obtain more information about unequal
distribution of transfer factor and ventilation. Techniques such as the intrabreath or three
equations method will probably be more important in the near future.
For patients who are not able to perform the single breath test or have a too small VC,
142
TRANSFER FACTOR FOR CARBON MONOXIDE
multiple breath diffusion tests have been developed. The rebreathing diffusion test is
recommended in these situations. Especially when measuring young children, a
rebreathing technique during rest ventilation is recommended.
The ECCS report [9] warns: "The association between TL,CO/VA and lung volume can
lead to difficulty in interpretation, particularly during childhood and adolescence, in
non-Caucasians and in patients in whom the total lung capacity is reduced". In patients
with restricted lung pathology the traditional comparison with predicted values for the
single breath diffusing capacity at predicted TLC is not correct. Dependent on the origin
of the restriction different ways are described to take the diminished alveolar volume into
account. This emphasises the importance of using an appropriate set of predicted values.
Further research on this particular issue is needed. The most advanced equipment may be
used to measure the transfer factor, but the results will be poor if the measured data are
not interpreted correctly!
Summary
The main function of the lungs is to establish exchange of O2 and CO2 between the
environment and the capillary blood. The gas transport across the alveolar-capillary
membrane can be measured by the transfer of carbon monoxide (CO). CO has a high
affinity for haemoglobin and is assumed to be absent in pulmonary capillary blood.
After inspiration, CO diffuses by the partial CO pressure gradient over the gas–blood
barrier from the alveoli into the capillary blood and disappears from the alveolar gas.
The decrease in CO fraction in the alveolar gas in a fixed time interval quantifies the
diffusing capacity of the lung. As not only diffusion but also chemical reactions affect
the CO transfer, the term "transfer" (T) rather than diffusion (D) is used.
Traditionally, gas transfer across the alveolo-capillary membrane is described in the
USA by the diffusing capacity for CO (DL,CO) and in Europe it is called the transfer
factor (TL,CO). However, DL,CO and TL,CO describe the same variable and are
interchangeable. Methods to determine the transfer factor TL,CO are the single
breath, the intrabreath and multiple breath methods. Each has its advantages and
limitations. The most important limitation of the single breath technique is the
required lung volume. Vital capacity (VC) has to be w1.5 L in order to obtain reliable
results. Traditional single breath measurements are inaccurate in the case of severe
airway obstruction due to inadequate time for equilibration of gases in the lung. Using
equipment that is based on fast responding gas analysers, conclusions of unequal
distribution of the diffusion characteristics may be drawn. A minimal VC of 1.5 L is
not required when using fast gas analysers. At reduced lung volume TL,CO/VA increases
and this may lead to erroneous interpretation of data in patients with a restrictive lung
disease. For the interpretation, it is important to take the possible influence of a
reduced VA or the influence of severe airway obstruction into consideration. In
patients who are not able to perform the single breath test and in small children the
transfer factor is determined with multiple breath methods. From the multiple breath
methods the rebreathing method is traditionally performed during hyperventilation.
Patients who are too ill to perform a single breath test, will also have problems with a
hyperventilation procedure. Therefore, a rebreathing method during normal, spontaneous ventilation was developed. When measuring the rebreathing transfer factor
during rest ventilation, it is important to realise that results are dependent on alveolar
ventilation and alveolar volume.
To minimise the variability in the diffusion measurement it is important to standardise
143
M. HORSTMAN ET AL.
these tests with respect to e.g. haemoglobin correction, body position, effect of O2 etc.
An important step forward is the use of European Respiratory Society/American
Thoracic Society guidelines.
Keywords: Intrabreath, multiple breath, obstructive disease, restrictive disease, single
breath, transfer factor.
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diffusing capacity at reduced alveolar volumes in children. Pediatr Pulmonol 1996; 21: 84–89.
Newth CJ, Cotton DJ, Nadel JA. Pulmonary diffusing capacity measured at multiple intervals
during a single exhalation in man. J Appl Physiol 1977; 43: 617–625.
Cotton DJ, Prabhu MB, Mink JT, Mink JT, Graham BL. Effect of ventilation inhomogeneity on
"intrabreath" measurements of diffusing capacity in normal subjects. J Appl Physiol 1993; 75: 927–
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Huang YCT, O’Brien SR, MacIntyre NR. Intrabreath diffusing capacity of the lung in healthy
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Filley CF, McIntosh DJ, Weight GW. Carbon monoxide uptake and pulmonary diffusing capacity
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Chinn DJ, Cotes JE, Flowers R, Marks AM, Reed JW. Transfer factor (diffusing capacity)
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145
CHAPTER 8
Clinical exercise testing
J. Roca, R. Rabinovich
Servei de Pneumologia i Allèrgia Respiratòria (ICT), Institut d’Investigacions Biomèdiques August Pi i
Sunyer (IDIBAPS), Hospital Clı́nic, Universitat de Barcelona, Barcelona, Spain.
Correspondence: J. Roca, Servei de Pneumologia, Hospital Clı́nic, Villarroel 170, Barcelona 08036,
Spain.
Impairment of exercise tolerance in chronic respiratory disorders, in particular chronic
obstructive pulmonary disease (COPD), has important implications on health-related
quality of life [1–3], hospitalisation rate [4, 5] and survival [6, 7]. Consequently, exercise
testing is progressively being considered an essential component in the routine clinical
assessment of these patients’ functional status.
Exercise intolerance results when a subject is unable to sustain a required work rate
sufficiently long for the successful completion of the task. The physiological cause, most
commonly, is an oxygen demand that exceeds the O2 conductance capability of the
oxygen transport chain. This is usually seen in physically fit individuals [8]. However, a
limited potential for oxygen utilisation at mitochondrial level must also be considered as
a factor of exercise limitation in healthy sedentary subjects [9–12]. The consequence of
exercise intolerance is a perception of limb fatigue, breathlessness or even, in some
conditions, frank pain. Exercise intolerance is the hall mark of range of cardiovascular,
respiratory and other systemic diseases, of which congestive heart failure (CHF) and
COPD are the most prominent.
Cardiopulmonary exercise testing (CPET) is a unique tool to assess the limits and
mechanisms of exercise tolerance. It also provides indices of the functional reserve of the
organ systems involved in the exercise response, with inferences for system limitation at
peak exercise. Moreover, CPET is useful for establishing the profiles and adequacy of the
system responses at submaximal exercise. Several studies [13, 14] have shown that the
functional reserve (i.e. aerobic capacity) of patients with COPD and interstitial lung
disease is not accurately predicted from resting lung function indexes.
The appropriateness of the integrated systemic responses are best studied utilising
incremental exercise testing, either as a ramp or small work-rate increments each of short
duration. CPET has been classically built around the rapid ramp-incremental exercise
test (performed on a cycle-ergometer or motorised treadmill), breath-by-breath
monitoring of cardiopulmonary variables (e.g. O2 uptake, CO2 output, ventilation,
heart rate) and formulation of graphical clusters of response profiles that optimise
estimation of key parameters, such as peak O2 uptake (V9O2) and the lactate threshold
and the characterisation of pertinent response profiles (e.g. V9O2–oxygen pulse, minute
ventilation–carbon dioxide production (V9E–V9CO2))
This provides a convenient means of: 1) determining whether the magnitude and
pattern of response of particular variables is normal with respect to other variables or to
work rate; 2) establishing a subject’s limiting or maximum attainable value for
physiological variables of interest; and 3) establishing exercise intensity domains, such as
the transition between moderate and heavy intensity exercise. It is important to
recognise, in this context, the difference between submaximal and maximal exercise
Eur Respir Mon, 2005, 31, 146–165. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
146
CLINICAL EXERCISE TESTING
levels. In submaximal exercise, the components of the O2 transport pathway can provide
adequate O2 flux between the air and the mitochondria. Mitochondrial oxidative
capacity has not been reached, symptoms are usually tolerable and muscle fatigue has not
occurred, or at least may be insufficient to impair performance appreciably. Figures 1
and 2 indicate the characteristics of different types of exercise protocols described in the
V 'O2 L·min-1
a)
4
b)
l
l
l
ll
ll
l
ll
l
ll l
ll
l l
l
l
ll
2
ll l
l
l
l l l l l l ll l l
l
l l ll l l l l
l
ll
l
ll ll l l l
l
l
l
l l l l l l l l l l l ll l l l l l l
l
l llll
0
l
l
l
l
l l
ll
l
l
l
l
l
l
l
l
l
Time
Time
Fig. 1. – Response of oxygen uptake (V9O2) to: a) a series of constant-work rate exercise tests, from moderate to
heavy exercise; and b) a ramp-incremental test. Note that peak oxygen uptake (??????) is not different between
the protocols (a and b). There was no evidence of plateau in oxygen uptake response (maximum O2 uptake).
Steady-state oxygen uptake was observed at moderate intensity constant-work rate exercise (a). Reproduced with
permission from [113].
2000
n
V 'O2 mL·min-1
1600
n
1200
s
l
s
s
s
800
n
l
l
l
l
l
l
l
s
l
l
s
l
l
l
l
l
l
l
l
400 n
l
0
0
1
2
3
4
6
7
5
Time min
8
9
10
11
Fig. 2. – Mean oxygen uptake (V9O2) profiles of the eight chronic obstructive pulmonary disease patients during
four different clinical exercise protocols (mean¡sem): incremental cycling (#); incremental shuttle ($); six-
minute walk test (+); and stairs climbing (%).
147
J. ROCA, R. RABINOVICH
text. At maximal exercise, symptoms have caused the patient to stop exercising. At this
stage, one or more of the following possibilities exist:
1) Limits to O2 transport have been reached and maximal V9O2 (V9O2,max) attained;
under such conditions, breathing 100% O2, for example, could increase V9O2,max [15].
2) Mitochondrial oxidative capacity has been reached and again the subject would be
considered to be at V9O2,max, but adding O2 would not raise V9O2.
3) Maximal exercise has occurred at a level that does not require maximal O2 transport
or maximal oxidative capacity and here exercise has been limited by unusually severe
symptoms. Under these conditions a plateau in O2 (V9O2,max) has not been reached and
the appropriate term is peak, rather than V9O2,max.
Subjects with lung disease often experience exercise intolerance at extremely low work
rates. There are many kinds of lung disease, however, and in any one patient the
structural and functional severity of the disease may range from the barely discernible to
the very severe. As a result, responses to exercise in patients with lung disease do not
show the tight stereotypical pattern of normal subjects. However, despite the widespread
clinical use of CPET, it does not typically provide a substantial improvement in primary
diagnostic power over the more classical clinical tools/assessments (e.g. spirometry).
What CPET can do, however, is: 1) reveal specific abnormalities that occur only when
support systems are stressed by exercise (e.g. dynamic hyperinflation in COPD); and 2)
provide a functional frame of reference for assessing the efficacy of interventions targeted
to ameliorate such abnormalities (e.g. bronchodilators for dynamic hyperinflation).
More recently, test paradigms designed to quantify endurance performance have evolved
and other exercise protocols, described below, have also become popular. It is of note,
however, that CPET is now highly developed in this regard and it is the accepted "gold
standard".
Responses to exercise in health and disease
There is nothing intrinsically different in the direction of the overall system response to
exercise comparing normal subjects and patients with lung diseases. Thus, as a patient
exercises harder, O2 consumption, CO2 production, ventilation and cardiac output all
increase to fulfill increased muscle bioenergetic requirements, as they do in the normal
subject [16], but peak levels attained are less, more so with increasing severity of the
disease. What may be different from normal in lung diseases regarding exercise responses
include: 1) resting function; 2) physical deconditioning; 3) the intensity and duration of
exercise that can be performed, and the relationship between intensity/duration and
symptom development; 4) the specific responses of the heart and lungs/chest wall to a
given exercise load in terms of rate, magnitude and performance limits; 5) the relative
importance of each part of the O2 and CO2 transport pathway in contributing to any
limitation of exercise that is found; 6) the relative importance of peripheral and
respiratory muscle fatigue; and 7) metabolic accompaniments of exercise, in particular
lactate release and accumulation, and high energy phosphate levels.
Pulmonary response to exercise in healthy subjects
It is well known that ventilation and cardiac output markedly increase during exercise
to match O2 transport with augmented cellular O2 requirements [17] (fig. 3). Since
ventilation increases to a higher relative extent than pulmonary blood flow, the ratio of
total alveolar ventilation to blood flow (overall V9A/Q9 ratio) rises rather substantially.
At moderate levels of exercise, the dispersion of the V9A/Q9 distributions does not change
148
CLINICAL EXERCISE TESTING
Pulmonary function
(ventilation and gas exchange)
Blood O2 carrying capacity
Cardiovascular system
(cardiac output and regional distribution of blood flow)
Muscle capillary O2 transfer capacity
Mitochondrial oxidative capacity
(cellular O2 utilisation)
Fig. 3. – Major elements of the O2 transport/O2 utilisation pathway. Integrated effects of all steps involved to
move oxygen from air to mitochondria are essential to determine the maximum capacity of the system. In
disease, nonuniformity of ventilation/perfusion ratios in the lung and/or metabolism/perfusion ratios in the
peripheral tissues may be of considerable importance.
[18–20] but the V9A/Q9 ratios at the mean of both ventilation and perfusion distributions
markedly increase due to the higher overall V9A/Q9 ratio. Consequently, the efficiency of
the lung as an O2 and CO2 exchanger improves at these exercise levels. Mixed venous
oxygen partial pressure falls dramatically during exercise because the relative increase in
V9O2 is considerably greater than that of cardiac output, and mixed venous carbon
dioxide partial pressure levels rise equally remarkably. Arterial PO2 levels generally
remain unchanged until extremely high levels of exercise are undertaken. Arterial PCO2
levels are also relatively stable until the appearance of high blood lactate levels generates
acidosis, even more ventilation, and thus a fall in Pa,CO2 levels. The alveolar–arterial O2
gradient (AaPO2), however, progressively increases with the level of exercise, reaching
values of 20–30 mmHg close to maximal exercise (VO2 peak) in average subjects, and
even greater (up to 40 mmHg or more) in some elite athletes [21]. Such an increase in
AaPO2 indicates inefficiency of pulmonary gas exchange during heavy exercise that is
even more apparent in other animal species, such as the horse [22]. It has been shown
that the increase in the AaPO2 during exercise is due, in part, to VA9/Q9 mismatching [18–
20] but it is mostly explained by alveolar-end capillary O2 diffusion limitation [19, 23].
Experimental studies suggest that development of subclinical pulmonary oedema [19, 24]
may explain the deterioration of pulmonary gas exchange during heavy exercise in elite
athletes.
Pulmonary response in lung diseases
In COPD patients, resting levels of V9E are abnormally high but, during exercise, the
slope between V9E and work rate is normal. For a given level of V9E during exercise, tidal
volume (VT) tends to be lower and respiratory rate (f) higher in patients than in healthy
subjects [25, 26]. Moreover, the O2 cost of breathing per unit ventilation is higher in
COPD patients than in healthy subjects. Impaired respiratory mechanics requires more
effort to move a given volume of air. Peak exercise VT is strongly related to vital capacity
in these patients [27]. They adopt two strategies during exercise to increase V9E [25]:
149
J. ROCA, R. RABINOVICH
4
3
2
Flow L·s-1
1
0
-1
-2
-3
-4
-5
0
1
2
3
Volume L
4
5
Fig. 4. – The resting maximal flow–volume curve from a chronic obstructive pulmonary disease patient is
represented by the solid line. The solid smallest loop corresponds to tidal volume at rest and the dashed curve
indicates tidal volume at maximal exercise. During exercise, end-inspiratory and end-expiratory lung volumes are
increased (dynamic hyperinflation) and expiratory flow limitation is seen over most of expiration. Reproduced
with permission from [113].
1) end-expiratory lung volume (EELV) increases, allowing higher maximum expiratory
flow rates (fig. 4). This dynamic hyperinflation does not occur in normal humans, who
show a fall in EELV during exercise [25]; and 2) inspiratory flow rate increases, so that
inspiratory time decreases and more time is available for expiration [25].
Impaired respiratory mechanics (dynamic hyperinflation) seems to play a major role
limiting exercise tolerance in these patients. During exercise in COPD, a balance is struck
between the need for ventilation and the high cost of breathing. The most common endresult is a small raise in arterial PCO2 and similar fall in Pa,O2. However, unless
pulmonary carbon monoxide transfer capacity (DL,CO) is severely impaired (v50%
predicted value), Pa,O2 does not fall during exercise, and may even increase in some
subjects. Studies using the multiple inert gas elimination technique in COPD show that
VA/Q mismatch is usually unaltered from that at rest, that shunts do not develop, and
that diffusion limitation also does not occur [28]. This is even the case when COPD is
severe [28]. In milder disease, there is evidence that small improvements in VA/Q
relationships may occur on exercise [29, 30], providing a partial reason for improvement
in arterial PO2. However, it is not infrequently observed that when the patient with
COPD is encouraged to maximal effort, sudden hypoxaemia and hypercapnia can
develop just before the patient quits exercising [28].
In a variety of chronic respiratory disorders, such as interstitial lung diseases (ILD)
and pulmonary vascular diseases (PVD), abnormally high resting levels of V9E and
normal slope between V9E and work rate during exercise are commonly observed, but not
dynamic hyperinflation, as seen in COPD patients. They do not change EELV
significantly during exercise [31]. Oxygen cost of breathing per unit ventilation is
increased in patients with ILD because the increased elastic recoil requires more
inspiratory muscle activity. They show a strong linear relationship between peak exercise
VT and vital capacity [31], suggesting that differences in peak VT are mainly due to
abnormal respiratory mechanics. During exercise, patients with ILD generally show
typical and substantial blood-gas changes, even at moderate effort. While arterial PCO2 is
150
CLINICAL EXERCISE TESTING
generally unaffected [31], Pa,O2 falls in almost all patients [32–34], sometimes severely, as
does mixed venous PO2. It is this profound degree of arterial hypoxaemia (and not
respiratory mechanics) that mostly limits exercise tolerance in ILD [35–38]. Worsening of
V9A/Q9 mismatching and shunt does not play a relevant role in exercise-induced
hypoxaemia seen in these patients [32]. Therefore, the blood gas changes on exercise are
mostly the consequence of: 1) insufficient increase of alveolar ventilation relative to the
rise in Pa,CO2; and 2) secondary effects from the fall in mixed venous PO2 causing a fall in
arterial PO2 [39]. Also, O2 diffusion limitation is seen in most ILD patients during
exercise, further adding to the hypoxaemia [32]. The presence of O2 diffusion limitation
in these patients despite the relatively low cardiac output at peak exercise (v10 L?min-1)
is likely related to the combination of: 1) an abnormally low mixed venous PO2; 2) a short
capillary transit time; and 3) some increased interstitial resistance for the diffusion of O2
from the alveolar gas to the capillary blood caused by the large collagen deposits there.
Exercise-induced hypoxaemia in patients with PVD has been found to be largely due to
the fall in mixed venous PO2, because there is no systematic change in V9A/Q9
relationships nor does diffusion limitation develop [32].
Haemodynamic responses to exercise in health and disease
In healthy subjects, cardiac output (Q9T) shows a linear increase in relation to O2
uptake during exercise. Likewise, both stroke volume and heart rate (HR) also increase
as V9O2 increases. In well-trained subjects, up to five-fold increase (y25 L?min-1) in Q9T at
peak exercise can often be seen. Systolic pulmonary pressure increases during exercise,
but pulmonary vascular resistance falls because of vascular recruitment. At systemic
levels, systolic pressure increases, but not diastolic pressure. It is of note, however, that
elite athletes at peak exercise show a potent sympathetic vasoconstriction at systemic
level inducing massive redistribution of cardiac output, which ensures preferential
perfusion to active skeletal muscle (due to local exercise-induced vasodilator effects)
while preserving blood flow and O2 delivery to essential organs such as the brain [40]. It
has been reported that in well-trained cyclists during maximal exercise, respiratory
muscles subvert blood flow that otherwise would have been directed to limb muscles. In
these subjects, unloading the respiratory system with proportional assist ventilation
resulted in an increase in both leg blood flow and leg vascular conductance [41, 42]. This
phenomenon is not seen in chronic respiratory patients because they are unable to reach
such extreme levels of O2 uptake during exercise but despite this they may show increased
O2 cost of breathing per unit ventilation [43].
In chronic respiratory diseases, pulmonary vascular abnormalities are present well
before frank heart failure occurs. There is pulmonary hypertension often even evident at
rest, and usually during exercise. The increase in pressure per unit increase in cardiac
output is some three times greater in these patients than in the normal subjects.
Contrary to the normal subjects in whom pulmonary vascular resistance normally falls
during exercise due to a combination of vascular recruitment and distension in the lungs,
in COPD, vascular resistance remains constant or may even rise. The vascular
destruction or obstruction that is well-known to occur in these diseases, together with
some distortion and also hypoxic vasoconstriction are the reasons underlying these
physiological abnormalities. Eventually, as the diseases progress, the right heart will
hypertrophy and ultimately fail, and clinically significant cor pulmonale will be present.
Despite the 2–3-fold increase in vascular resistance and high pulmonary artery pressures,
it is remarkable that even in advanced lung disease the heart can pump essentially
normally as a function of filling pressure, as shown from the limited data available.
At peak exercise, systemic O2 delivery is clearly below normal level [25]. While the
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obvious culprit is impaired pulmonary function, it is not always through a reduction in
oxygen saturation in arterial blood that systemic O2 delivery is primarily reduced, since
despite V9A/Q9 inequality and reduced effective alveolar ventilation, hypoxaemia may
not necessarily provoke a marked fall in arterial O2 content [25]. It is well accepted that
cardiac output at peak exercise is always well below normal levels. However, in COPD
patients as in normal subjects, cardiac output increases linearly in relation to oxygen
uptake as work rate increases during incremental exercise, such that cardiac output at a
given submaximal O2 uptake [44] is close to the expected normal value. It should be
noted, however, that the rise in cardiac output during exercise is usually achieved
through a higher heart rate and lower stroke volume than in healthy subjects.
Since total ventilation, cardiac output and exercise intensity remain closely coupled in
COPD as in health, the inability to raise ventilation appears as the principal governor of
the O2 transport process: a low ceiling on ventilation means a low ceiling on cardiac
output and thus on systemic O2 delivery. It should be mentioned that the mechanisms
that couple ventilation to cardiac output during exercise are still not well understood.
Montes de Oca et al. [45] proposed that the large pleural pressure swings observed
during exercise can be paramount to constrain left ventricular function, thus limiting
both peak cardiac output and exercise tolerance in very severe COPD patients. The
coupling between whole-body O2 uptake and cardiac output during exercise implies that
the O2 difference between arterial and mixed venous blood and the fractional O2
extraction are normal or near normal [11, 46]. The cardiac response to exercise in patients
with ILD is similar to the description for COPD patients. In contrast, patients with PVD
show different cardiac response to exercise. Certainly, at peak exercise cardiac output is
lower. More importantly, however, the slope of the relationship between V9O2 and
cardiac output appears different. This suggests that for any given degree of exercise (i.e.
V9O2), cardiac output in patients with PVD does not increase as much as in controls or
patients with COPD or ILD. This abnormal behaviour is likely related to the increased
after-load of the right ventricle [47–49]. As expected, patients with PVD have, at rest,
pulmonary artery hypertension and increased pulmonary vascular resistance. Compared
with patients with COPD and ILD, patients with PVD show, by far, the worse
haemodynamic situation. During exercise, pulmonary artery pressure increases in direct
proportion to the increase in cardiac output and reaches extremely high values. This
indicates the lack of pulmonary vascular reserve. In fact, the pathologically elevated
pulmonary vascular resistance seen at rest does not change substantially during exercise.
Muscle oxygen utilisation in health and disease
It has been reported that well-trained males show O2 supply dependency of maximum
O2 uptake [8], indicating that mitochondrial capacity does not constitute the rate limiting
factor for maximum exercise performance. In contrast, data from healthy sedentary
subjects [9, 10] strongly suggest that muscle mitochondrial function is a limiting step for
maximum O2 uptake (fig. 5). Studies including direct measurements of cell PO2
saturation during exercise, breathing different inspiratory oxygen fraction (FI,O2),
further indicates that sedentary subjects do not show O2 supply dependency of V9O2,max
[12]. The plasticity of skeletal muscle during a high-intensity physical training
programme [50] fully accounts for the differences alluded to between athletes and
sedentary subjects. The scenario is far more complex in patients with COPD. Femoral
blood flow (Qleg) measurements in patients with moderate-to-severe airflow limitation
[11, 51] have shown, as for cardiac output, a marked reduction in leg blood flow at peak
exercise. However, leg blood flow (and leg O2 delivery) [11, 52] at a given submaximal
whole-body O2 uptake (and leg V9O2) is above normal, which may indicate increased
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CLINICAL EXERCISE TESTING
Maximum V 'O2,leg L·min-1
0.8
1.0
0.6
s
0.13
s
s
n
0.4
n
0.21
n
1.0
0.21
0.13
0.2
0.0
0
400
Peak power g
800
1200
Fig. 5. – Quadriceps maximal oxygen uptake (V9O2,leg) (y-axis) plotted against maximum work rate (x-axis)
(mean¡sem) in healthy subjects (') and in chronic renal patients (&) breathing 13%, 21% and 100% inspired
O2 concentrations (inspiratory oxygen fraction (FI,O2) 0.13, 0.21 and 1.0, respectively). While chronic renal
patients increased V9O2,max and maximum work rate (Wmax) proportionally to the FI,O2 increase, indicating
O2 supply dependency of V9O2,max, healthy sedentary subjects did not show any relationship between exercise
performance and changes in O2 transport (and in cell oxygenation) suggesting that mitochondrial capacity,
but not O2 transport, was limiting V9O2,max. Reproduced with permission from [12].
peripheral muscle O2 demand. Moreover, poor muscle capillary network in these patients
[53] seem to suggest that low peripheral O2 diffusion capacity may also contribute to
exercise-induced cell hypoxia, even in the absence of arterial hypoxaemia. Increased
lactate production [11, 54–56] is responsible for the fall in muscle pH, which, in turn, may
play a role in determining exercise intolerance in these patients [56]. Premature lactic
acidosis during exercise in COPD patients has been associated with reduced oxidative
enzyme concentrations in the lower limb muscles [54, 55] that can be, at least partly,
reversed by physical training.
Several studies [11, 57, 58] exercising different muscle groups in heterogeneous groups
of COPD patients have consistently shown lower cellular bioenergetic status (31PNuclear magnetic resonance spectroscopy) and lower pHi than those seen in healthy
sedentary controls at equivalent levels of exercise. There is evidence [11] suggesting that
muscle deconditioning plays a major role to explain the disturbances of skeletal muscle
bioenergetics in COPD patients. Recent lines of evidence indicate that intrinsic skeletal
muscle dysfunction may be present in patients with COPD, as well as in other chronic
disorders, such as CHF [59–61]. Abnormal redox status [59–62] plays a central role
prompting muscle mass wasting particularly in susceptible subsets of COPD patients.
Factors determining exercise performance: integrated response
It is presently well accepted that the level of exercise tolerance is set by the integrity of
each of the functions involved in the O2 transport/O2 utilisation system, as well as by
proper interactions among all of the physiological responses alluded to above [63].
Complex integrative pathways both at whole body level and at cellular level have been
identified. Since not only intracellular pH [64], but also cell PO2 [65] has been shown to
modulate mitochondrial function, O2 transport (cell PO2) and O2 utilisation
(mitochondrial capacity) cannot be analysed as separate systems.
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J. ROCA, R. RABINOVICH
Also of major interest are the events surrounding peak or maximal V9O2 and the
physiological basis of why peak or maximal V9O2 is reduced as it almost always is in disease.
In this regard, it must be noted that the amount of V9O2 achieved by a given patient is not
only set by the intrinsic characteristics of the system, it also depends on several other factors
that modulate the physiological response of the whole body, such as: 1) environmental
conditions (altitude above sea level, FI,O2); 2) amount of exercising muscle mass (cycling,
walking, localised quadriceps exercise); and 3) type of exercise protocol (incremental,
endurance test, 6-min walking distance test (6MWT), shuttle test, etc.) (fig. 2). Since the
catabolic capacity of the myosin ATPase is such that it outstrips by far the capacity of the
respiratory system to deliver energy aerobically, exercise tolerance (V9O2,max) is determined
by the capacity of the O2 transport/O2 utilisation system rather than by the muscle’s
contractile machinery. Two physiological muscle properties (muscle strength and muscle
fatigability) may modulate functional performance of the patient in daily life activities, as
well as during clinical exercise testing. Muscle strength is defined as the force generated by a
muscle. It is determined by the number and type of motor units recruited; whereas muscle
fatigue has been defined as a loss of contractile functions (force, velocity, power or work)
that is caused by prolonged exercise and is reversible by rest. Factors involved in muscle
fatigue are complex, mainly: 1) contractile machinery; 2) muscle respiratory capacity; and
3) redox status of the muscle. In practical terms, it may be useful to consider two different
scenarios (V9O2,peak and V9O2max) (fig. 1). These are the following:
1. A peak V9O2 has been reached without evidence of V9O2 plateauing. This is perhaps
the commonest outcome in the clinical setting. Taken as it is, one cannot say
whether this peak V9O2 is limited by O2 supply, mitochondrial oxidative capacity, or
perhaps neither (i.e. symptoms are so severe that neither O2 supply nor
mitochondrial function have been fully exploited). In these circumstances, it will
be useful to identify the V9O2 at which the transition from moderate to heavy
exercise took place (lactate threshold) and evaluate the organ system responses
(ventilation, gas exchange, heart rate, etc.) during submaximal exercise and at peak
V9O2. Despite not having information about the capacity of the system (a plateau of
O2 uptake was not identified), we will know about: 1) the physiological burden
imposed by exercise; and 2) the reserve of the system depending upon the location
of the transition from moderate to heavy exercise.
2. A plateau in V9O2 at maximal exercise is clearly identified such that the subject
achieved his/her maximum exercise (maximum O2 uptake) capacity in that
particular setting or there is physiological evidence that they are very close to
maximum. In this circumstance, two situations may be faced:
2.1. V9O2,max is the result of having reached mitochondrial oxidative capacity. In this
scenario, the key concept is that acute increases in O2 supply to the mitochondria
would not lead to any further increase in V9O2,max. In other words, no O2 supply
dependency is observed by giving 100% O2 to breath or by blood transfusion.
2.2. V9O2,max is the result of having reached limits to the supply of O2. In this
circumstance, one or more components of the integrated O2 transport system
(the lungs, heart and blood vessels, blood and muscles) has reached maximal
capacity for the given conditions and it can be tested experimentally by
augmenting any one of the components alluded to above.
Clinical indications and exercise protocols
There is a range of indications for CPET. It is useful, for example, in the diagnosis of a
range of disease conditions, namely: exercise-induced asthma, cardiac ischaemia,
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foramen ovale patency with development of right-to-left shunt during exercise, and
McArdle’s syndrome [66]. In addition, CPET provides information on dysfunction,
monitoring or prognostic value in a wide range of conditions. However, an adequate
identification of the clinical problem requiring study should be considered a necessary
prelude to CPET, as should an appropriate assessment of the patient by: 1) medical history;
2) physical examination; 3) chest radiograph; 4) pulmonary function testing; and 5)
electrocardiogram (ECG). The clinical problem that prompts the CPET and the specific
aims of the test (i.e. assessment of exercise tolerance, analysis of pulmonary gas exchange
during exercise, etc.) determine both the type of exercise protocol to be used and the
variables to be considered in the interpretation of the test. Assessment of exercise tolerance
and potential limiting factors constitutes the most important indication of CPET. This is
particularly important to evaluate dyspnoea, but also to assess the degree of impairment in
several chronic diseases. Appropriate use of CPET allows the investigator: 1) to quantify
the degree of abnormal limitation and to discriminate among causes of exercise intolerance;
2) to differentiate between dyspnoea of cardiac or pulmonary origin when respiratory and
cardiac diseases co-exist; and 3) to analyse unexplained dyspnoea when initial pulmonary
function impairment does not provide conclusive results.
A second area of indication of CPET is pre-operative assessment in different
conditions, namely, major abdominal surgery in elderly patients [67, 68]. Also, CPET are
indicated in lung cancer resectional surgery and lung volume reduction surgery.
Information on predicted post-operative lung function: 1) helps to modulate the amount
of lung parenchyma to be resected; and 2) determines the type of peri-operative strategy
needed to prevent post-surgical complications. Resting pulmonary function tests are
considered adequate to evaluate patients with low risk (forced expiratory volume in one
second w2 L and DL,CO within the reference limits) of post-surgical complications [69–
74]. However, CPET play a pivotal role in the evaluation of patients with moderate-tohigh risk [72, 73, 75, 76]. Assessment of patients included in transplantation programmes
(lung, heart) also constitutes an indication for CPET.
CPET should always play a central role assessing candidates before the rehabilitation
programme and in the subsequent modulation of the exercise prescription, whereas
simpler tests (i.e. 6MWT) are useful for monitoring during the rehabilitation programme.
Finally, assessment of impairment-disability also constitutes a central indication of
CPET. It is now well accepted that CPET provides different and relevant information in
impairment-disability evaluation [77–79], compared to resting cardiopulmonary
measurements [80]. Consequently, CPET constitutes a key tool in this area.
Exercise protocols
The goal of CPET protocols is to stress the organ systems involved in the exercise
response in a controlled manner. For this reason the testing generally involves exercising
large muscle groups, usually the lower extremity muscles. A key requirement is that
exercise stimulus must be quantifiable in terms of the external work and power
performed. The appropriateness of the integrated systemic responses to the tolerable
range of work rates is best studied utilising incremental exercise testing. This provides a
smooth incremental stress to the subjects so that the entire range of exercise intensities
can be spanned in a short period of time. The recommended incremental exercise testing
protocol, usually electronically-braked cycle ergometry with constant pedalling
frequency, of 60 rpm is recommended. Equivalent results are obtained when work
rate is either increased continuously (ramp test) or by a uniform amount each minute (1min incremental test) until the patient is limited by symptoms (he/she cannot cycle
w40 rpm) or is not able to continue safely. The increment size should be set according to
155
J. ROCA, R. RABINOVICH
the characteristics of the patient in order to obtain y10 min duration of the incremental
part of the protocol. This may represent incremental rates of 10–20 W per minute in a
healthy sedentary subject or less in a patient. Sufficient density of data to be acquired in a
test lasting v20 min from start to finish, including: 1) measurements at rest; 2) 3 min of
unloaded exercise; 3) incremental exercise (y10 min); and 4) 2 min recovery, at least.
Standard noninvasive CPET carried out whilst breathing room air (FI,O2=0.21) involves
acquisition of breath-by-breath expired O2 and CO2 concentrations (expiratory oxygen
fraction and expiratory carbon dioxide fraction, respectively), work rate, expired airflow,
HR and systemic arterial pressure as primary variables. ECG and pulse oximetry should
be continuously monitored during the test. It is useful to establish a sense of the patient’s
exercise-related perceptions during the exercise test and at the point when the subject
discontinues exercise. This includes exertion, dyspnoea, chest-pain and skeletal muscle
effort. Quantifying these perceptions should be done using standardised rating
procedures (Borg scale, visual analogue scale (VAS) etc.).
Proper evaluation of pulmonary gas exchange in patients with lung disease requires
assessment of arterial respiratory blood gases [81]. In these cases, arterial cannulation
(preferentially radial, or brachial) is needed (Pa,O2 and Pa,CO2 measurements and
calculation of AaPO2) [81, 82]. This also provides information on acid-base status (pH,
Pa,CO2 and base excess) and allows continuous monitoring of systemic arterial blood
pressure during the test. However, while "arterialised venous blood" (e.g. from the dorsum
of the heated hand) gives good values for PCO2 and pH it is not appropriate for PO2.
Furthermore, estimation of arterial respiratory blood gases through expired O2 and CO2
profiles or "transcutaneous" electrodes and pulse oximetry should not be used as indices of
arterial PO2 and PCO2 during exercise [83–85]. It is important to recognise that arterial
blood sampling immediately after exercise does not provide an adequate assessment of
blood gas values at peak exercise. However, while pulse oximetry does not indicate arterial
PO2, it does provide valuable information on oxyhaemoglobin saturation during exercise.
If the ergometer used in the CPET is a motor driven treadmill, then the Balke’s
protocol [81, 86] is considered the most appropriate for its simplicity. The speed of the
treadmill is kept constant (3–3.5 mph) during the protocol while the slope is
progressively increased (1–2% min-1). It is of note, however, that the assessment of
the relationships between oxygen uptake and external work rate is more accurately
carried out using a cyclo ergometer than using a treadmill.
Alternative protocols can be considered for specific purposes [87]. Simpler tests, such
as step tests or timed distance walks (i.e. 6MWT or 12 min-walk) are widely used and
they can provide measures of exercise tolerance but are not as useful in diagnosis as
incremental tests [88–90]. The timed walking tests have been extensively used in the
clinical evaluation of patients with chronic cardiopulmonary disorders mainly because of
their simplicity. A present, these tests are recognised to add prognostic information
useful to the staging of patients with COPD [4, 7], primary pulmonary hypertension [91]
and congestive heart failure [92]. Timed walking tests have shown to be sensitive to
changes after interventions such as inhaled bronchodilators [93], volume reduction
surgery [94] and pulmonary rehabilitation [95, 96]. The 6MWT, for example, is currently
performed in a large number of rehabilitation programmes. Recent studies [97] suggest
that encouraged 6MWT is a strenuous protocol that evaluates sustainable exercise
performance; that is critical power. The 6MWT and the incremental cycling protocols
should be considered complementary tests.
Constant-work rate protocols can result in steady-state responses when work rate is of
moderate intensity. In contrast, constant work rate of high intensity for the individual
typically results in continually changing values in most variables of interest. Consequently,
attainment of, or failure to attain, a steady-state V9O2 during a constant-load test can be used
to determine if a particular task is sustainable by the individual. During a constant-work
156
CLINICAL EXERCISE TESTING
rate protocol, the period of dynamic adjustment to a constant-work rate test provides
information regarding the dynamic behaviour of lung function, haemodynamics and tissue
O2 utilisation. However, there is to date virtually no information on the confidence limits,
reproducibility and predictive value of the derived parameters in patient populations.
Consequently, the utility of quantifying dynamic responses to constant-work rate exercise in
clinical exercise testing remains to be established. The constant-work rate protocols are,
however, useful to assess the impact of a given intervention on the system responses to
exercise (i.e. bronchodilator therapy) [98]. Alternatively, the use of high intensity constantwork rate to assess exercise-induced asthma has been traditionally used in the clinical
setting, but it might be progressively substituted [99].
Testing procedures
Cardiopulmonary exercise testing should be conducted only by adequately trained
personnel with a basic knowledge of exercise physiology. Technicians familiar with normal
and abnormal responses during exercise and trained in cardiopulmonary resuscitation
(CPR) should be present throughout the test. CPET should be performed under the
supervision of a physician who is appropriately trained to conduct exercise tests and in
advanced CPR. The degree of subject supervision needed during the test can be determined
by the clinical status of the subject being tested and the type of exercise protocol. While it is
preferable for the physician to be present during the test, if not he/she must be readily
available to respond as needed. Additional roles for the physician are the evaluation of the
patient immediately before the test and the interpretation of the results.
Patient preparation
At the time of scheduling, the subject should be instructed to adhere to his/her usual
medical regimen; he/she should not to eat for at least 2 h before the test, avoid cigarette
smoking and caffeine, and dress appropriately for the exercise test. A brief history (with
detailed inquiries about the medications) and physical examination should be done to
rule out contraindications to testing. Results of recent resting pulmonary function tests,
as a minimum forced spirometry, should be available for patients in whom pulmonary
disease is suspected.
On arrival at the CPET laboratory, a detailed explanation of the testing procedure and
equipment should be given to the patient outlining risks and potential complications as
described below. The subject should be told how to perform the exercise test and the
testing procedure should be demonstrated if needed. The patient should be encouraged to
ask questions to reduce any anxiety. The patient needs to become familiar with the
equipment. If the treadmill is used, time is provided for several practice trials of starting
and stopping until the patient feels confident. If the cycle ergometer is used, the seat
height is adjusted so that the subject’s legs are almost completely extended when the
pedals are at the lowest point and the cycling rhythm practiced. Before the test, the ECG
electrodes are carefully placed and secured after preparing the skin to ensure good
recordings (if necessary, the area of the electrodes placement should be shaved). A
sphygmomanometer cuff is placed on the upper arm. The mouthpiece and noseclip are
then tried and the position adjusted until adopting a comfortable position. The patient is
informed that it is acceptable to swallow with the mouthpiece in place and that he/she
must signal any unexpected difficulty by the signal "thumbs down". The patient is advised
to point to the site of discomfort if chest or leg pain is experienced.
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J. ROCA, R. RABINOVICH
During the test, the patient is encouraged to carry on with a regular pedalling cadence.
Symptoms and degree of discomfort are periodically checked (see below safety
precautions). Good communication with the patient throughout the whole procedure
increases the subject’s confidence and predisposes to good effort. During recovery, the
patient is told to continue to pedal, without external work load (or walk at a slow pace on
the treadmill), for at least 2 min during recovery in order to prevent fainting and to
accelerate lactate removal. At the point when the subject discontinues exercise, after
removal of the mouthpiece, the physician should ask for symptoms (type and intensity)
that prompted the patient to stop exercise. If blood gas analysis is done, a last blood
sample is taken at 2 min of recovery. If the test does not provide adequate diagnostic
information because of premature termination or inadequate cooperation of the patient,
it should be repeated after a resting period of 30–45 min.
Although CPET may be considered to be a safe procedure, risks and complications
have been reported. Good clinical judgment should be paramount in defining indications
and contraindications for exercise testing [100]. Cardiac (bradyarrhythmias, ventricular
tachycardia, myocardial infarction, heart failure, hypotension and shock) and
noncardiac (musculoskeletal trauma, severe fatigue, dizziness, fainting, body aches)
complications of CPET have been reported. Consequently, during the test, the personnel
should be alert to any abnormal event. The indications to stop the test must be clearly
established and known by all the personnel involved in testing. These indications include
symptoms such as: 1) acute chest pain, 2) sudden pallor, 3) loss of coordination, 4)
mental confusion, and 5) extreme dyspnoea; and signs such as: 1) depression of ST
segment w0.1 mV (less specific in females), 2) T-wave inversion, 3) sustained ventricular
tachycardia, and 4) fall in systolic pressure either below the resting value or y20 mmHg
below its highest value during exercise testing. Relative indications to stop the test are: 1)
polymorphic and/or frequent premature ventricular beats; and, 2) hypertension
(w250 mmHg systolic, w130 mmHg diastolic). If the exercise test has been stopped for
one of the above-listed reasons, the patient should be monitored in the CPET laboratory
until symptoms or ECG modifications have completely cleared. Admission to hospital
for longer observation or more often for complementary investigation will be necessary
in very rare cases. If necessary, intensive care can be administered on site. Full
cardiopulmonary resuscitation equipment should be available in the CPET laboratory.
Interpretation strategies
The greatest diagnostic potential and impact on the clinical decision making process of
exercise testing should rely not on the utility of any one individual measurement,
although some are obviously more important than others, but rather on their integrated
use. Identification of a cluster of responses characteristic of different diseases is often
useful. The major portion of the interpretation strategy is focused on CPET results
generated during maximal, symptom-limited, incremental exercise testing. This is
currently the most popular, albeit not the exclusive protocol. Often, insufficient attention
is paid to trending phenomena as the work rate progresses from submaximal to peak
levels. To facilitate this type of analysis, the results should be formatted in an appropriate
manner. Figure 6 displays data obtained in a normal subject performing cycle ergometry,
using an ergometer that utilises an "assist" to provide an actual zero-watt work rate at
"unloaded" pedalling. Figures 6a–d provide, in addition to the peak V9O2, the variables
commonly used to provide an indirect estimation of the lactate threshold. That is,
identification of the O2 uptake at which the transition between moderate to heavyintensity exercise occurs. Figure 6e (O2 uptake versus work rate) reflects the exercise
158
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e)
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2.5
n
n
n
n n n
n n
n
n n
nn
n nn n
nnnn n nn
n n
n
nn
n
n
n
n
0
n
nn
nn
nnnn
nnnn
n n
nn
nn
nn n
n
nn
n
n n
n
n nn
n n nn
n
nn
nnn nnn
n
n
n
50
100 150
Work rate W
200
250
f) 150
n
125
100
V 'E L·min-1
55
50
45
40
35
30
25
20
0.0
2.0
1.5
n
nnn
nn
n
nn
*
2.5
n
75
50
25
0
0.0
g) 180
160
140
120
100
80
0.0
nnn n
nn nn
nn n
nnnn n
n
n
n
nn
nn
nnnn
nnnn
n
nn n
n
n
n
n
0.5
nn
n
nn
nnn
n nn
nnn
n
nn
n
nn n
n
n
nn
n
nn
n
n nn
1.0
1.5
2.0
V 'CO2 L·min-1
n
nn
n
n n
n nn
n
nn
n
nn
nnn n
nn n nnn n
nn
n
n
n
nn nn
nn
n
n nn nn
n
nn
nnn
nn
nnn
n n n
n nnnn
n
nn
nn
n nn
n
n
nn
nn
n n
n
n nn
nnn
nn n
nn n
n
n
n
n n
nnnn
nn n
nn
n
nn
n
nn
n
nn
n
n
nn
n
nn n
n
n
n nn
n
nn
n
nnn
n
nnnn
n
n
nnn nn
nn nn nnn
n
nnnnn
nn
nnnn
n
n
n
0.5
1.0
1.5
V 'O2 L·min-1
2.0
2.5
* 16
14
12
10
8
6
4
2
2.5
h) 2.5
2.0
n
n
n nn
nnn
n nn
n
n n n
0.0
n nn n
n nn
n
nn
n n
n n
n
n nnn
nnnnnnnn
n
nn
n n nnnnn n
n nnnnnnnnn n nn
nn
n
nnnn
n
nn nn
n
n
0.5
1.0
1.5
V 'O2 L·min-1
n
n
n
n
n
n
n
n
n
nn
n
nn
n n
1.5
n
n
0.5
2.0
0.0
2.5
n
n nn nn n
n
nn n
n
n n
n n
nn
n
nn
n
n
n n n
n
nn n
n
n
nnn
n
nn
n n n
n
n
n
n
nn
n nn
nn
nnn nn n n
nnn nnn
n
n
n
nnn
nn
nn
n
n
n
nn
1.0
n
n
n
VT L
V 'E/V 'O2
b)
nn
nnnn
nn
nn
nn
nn
nn n
n
fc beats·min-1
0.0
0.0
nn
n
nn
nnn
nn
n
nn
n
V 'E/V 'CO2
0.5
n
nn
n
n
nn
n
V 'O2/fc mL·beats-1
2.0
nn
n
nn
n n
n
nn
nn
n
n
V 'O2 L·min-1
2.5
P ET,CO2 mmHg
V 'CO2 L·min-1
a)
0
25
50
75 100
V 'E L·min-1
125
150
Fig. 6. – Exercise performance in a healthy sedentary male subject. The basic plots for the interpretation of
cardiopulmonary exercise testing are reported. In plots a–d, in addition to peak oxygen uptake (V9O2), the
variables commonly used to indirectly estimate lactic threshold (LT) are given. That is, the V9O2 at which the
transition from moderate to high intensity exercise occurs is identified (vertical dashed line). The expected LT
for a healthy subject (55% of predicted V9O2,peak) is indicated in plots a) to d) by a small arrow (continuous
line). Predicted V9O2,peak is indicated in a) by an arrow (dashed line). In plot e), V9O2 versus work rate reflects
the exercise efficiency and limits of exercise tolerance of the subject; with the expected peak exercise performance
represented by the asterisk. Plots f) and h) indicate minute ventilation (V9E) versus carbon dioxide uptake
(V9CO2) and tidal volume (VT) versus V9E, respectively; these two plots describe the characteristics of the
ventilatory response during submaximal and peak exercise. Finally, plot g) presents characteristics of the
haemodynamic response to exercise with estimated peak heart rate represented by the asterisk and predicted
peak O2 pulse by the arrow. PET,O2: end-tidal pressure of oxygen; PET,CO2: end-tidal pressure of carbon dioxide;
RER: respiratory exchange ratio. Reproduced with permission from [113].
159
J. ROCA, R. RABINOVICH
efficiency and the limits of exercise tolerance of the subject. Figure 6f (ventilation versus
CO2 output) and figure 6h (tidal volume versus ventilation) characterise aspects of the
ventilatory response during submaximal and maximal exercise. However, some
investigators find the relationship between V9E and V9O2 during such tests to be
useful. Finally, figure 6g, which plots heart rate (and O2 pulse) versus O2 uptake, is
informative with respect to the characteristics of the haemodynamic response to exercise.
The next step is to choose adequate reference values to establish patterns of normal or
abnormal response. Available reference values and present limitations in this particular
issue are discussed below. Relatively few studies have evaluated the sensitivity, specificity
and predictive value of patterns of measurements in distinguishing among different
clinical entities. Even more importantly, the precise role of clusters of variables
commonly used in the decision making process in well identified diseases (i.e. evaluation
of ILD, pre-operative evaluation for resection lung cancer surgery, etc.) is insufficiently
known. For the future, studies addressing the use of likelihood ratios might be even more
useful to clinicians than sensitivity/specificity, since likelihood ratios refer to actual test
results before disease status is known. This shift to an evidence-based approach for
CPET interpretation will hopefully provide important answers to clinically relevant
questions that are not immediately available.
Selection of appropriate reference values is an important step to establish patterns of
normal or abnormal response to exercise stress. An initial analysis of available data on
healthy subjects [88, 89, 101–111] clearly indicated that only some of these studies [89,
103, 105–107] fulfil minimum requirements to be considered as candidates to be used in
the clinical setting. Blackie et al. [105] cover a limited age span (from 55–80 yrs) and
Bruce et al. [107] provide data obtained with treadmill in a population of rather
physically fit people. Hence, the analysis of potential studies in healthy sedentary people,
providing prediction equations for V9O2,peak obtained with cycling incremental exercise
testing, is then even more reduced to three sets [102, 106, 112]. Reference values estimated
by fairbarn et al. [106] are consistently higher than those provided by Jones et al. [102],
both in males and females. Predicted values by Wasserman et al. [89] and Hansen et al.
[112] are closer either to Jones et al. [102] or to Fairbarn et al. [106], depending upon the
values of height-weight of the subject in whom the equations are used. The characteristics
of the presently available prediction equations for peak O2 uptake (and peak work rate)
clearly impose limitations to the interpretative strategy. Moreover, except for HR in the
study of Fairbarn et al. [106] the profile of response in healthy sedentary subjects (i.e.
from submaximal to peak exercise results) are not available. Further, adequate
prediction equations for the most important variables obtained from the same group of
reference subjects are not currently available.
Summary
The role of the O2 transport/O2 utilisation system determining maximum O2 uptake
has been analysed in an integrative manner. The system responses to exercise in
healthy subjects (athletes and sedentary) and in common pulmonary diseases have
been examined. Finally, basic principles of exercise testing and interpretation of the
results have been reviewed.
Keywords: Aerobic capacity, cardiopulmonary exercise testing, endurance, lactic
threshold, oxygen uptake, work rate.
160
CLINICAL EXERCISE TESTING
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and ranges for ventilation and breathing pattern at maximal exercise. Chest 1991; 100: 136–142.
Blackie SP, Fairbarn MS, McElvaney NG, Morrison NJ, Wilcox PG, Pardy RL. Prediction of
maximal oxygen update and power during cycle ergometry in subjects older than 55 years of age.
Am Rev Respir Dis 1989; 139: 1424–1429.
Fairbarn MS, Blackie SP, McElvaney NG, Wiggs BR, Paré PD, Pardy RL. Prediction of heart
rate and oxygen uptake during incremental and maximal exercise in healthy adults. Chest 1994;
105: 1365–1369.
Bruce RA, Kusumi MS, Hosmer D. Maximal oxygen intake and nomographic assessment of
functional aerobic impairment in cardiovascular disease. Am Heart J 1973; 85: 546–562.
Vogel JA, Patton JF, Mello RP, Daniels WL. An analysis of aerobic capacity in a large United
States population. J Appl Physiol 1986; 60: 494–500.
Pollock ML, Bohannon RL, Cooper KHE. A comparative analysis of four protocols for maximal
treadmill stress testing. Am Heart J 1976; 92: 39–46.
Shephard RJ, Allen C, Benade AJ. The maximal oxygen intake - an international reference
standard of cardiorespiratory fitness. Bull World Health Organization 1968; 38: 757–764.
Taylor HL, Buskirk E, Henschel A. Maximal oxygen intake as an objective measure of cardiorespiratory performance. J Appl Physiol 1955; 8: 73–80.
Hansen JE, Sue DY, Wasserman K. Predicted values for clinical exercise testing. Am Rev Respir
Dis 1984; 129: S49–S50.
Roca J, Whipp BJ, eds. Clinical Exercise Testing. Eur Respir Mon 1997; 2 (no. 6).
165
CHAPTER 9
Respiratory function measurements in
infants and children
P.J.F.M. Merkus*, J.C. de Jongste*, J. Stocks#
*Division of Respiratory Medicine, Dept of Paediatrics, Sophia Children’s Hospital – Erasmus Medical
Centre, Rotterdam, the Netherlands. #Portex Anaesthesia, Intensive Therapy and Respiratory Medicine
Unit, Institute of Child Health, London, UK.
Correspondence: P.J.F.M. Merkus, Division of Respiratory Medicine, Dept of Paediatrics, Sophia
Children’s Hospital – Erasmus Medical Centre, PO Box 2060, 3000 CB Rotterdam, the Netherlands.
Most children above the age of 7–8 yrs can perform the full range of tests available for
older individuals, using similar protocols to those described elsewhere in this
Monograph. By contrast, assessments in young children and infants have generally
been restricted to specialised research establishments, due to the lack of suitable
equipment and the complexity of undertaking such measurements. The realisation that
insults to the developing lung may have life-long effects and that much of the burden of
respiratory disease in childhood and later life has its origins in infancy and early
childhood has emphasised the need to develop and standardise sensitive methods of
assessing respiratory function in infants and young children [1, 2]. During the past few
years there have been concentrated efforts to improve the feasibility of assessing lung
function in preschool children. With specially trained operators and a suitable
environment, many pulmonary function tests (PFTs) now appear to be feasible in at
least 50% of 3-yr-old children and in the majority of children aged w4 yrs.
The aims of this chapter are to provide an overview of:
. the differences in assessing lung function in infants and preschool children
compared with older cooperative subjects;
. which tests are feasible in infants and young children;
. the limitations of applying these tests; and
. the problems associated with interpreting results in this age group.
Throughout this chapter, the focus will be on the most widely used pulmonary
function tests for this age group. For simplicity, the term "infant PFTs" will refer to
measurements in sleeping infants and young children (agedv2 yrs), whereas "preschool"
will apply to those tests used in awake young children (aged 3–6 yrs).
Assessing lung function in different age groups
Infants and toddlers below 2 yrs of age
Marked developmental changes in respiratory physiology occur during the first years
of life. The major issues in undertaking PFTs in children agedv2 yrs relate to sleep state,
sedation, ethical issues, posture and the need to miniaturise and adapt equipment for
measurements in small subjects who cannot cooperate actively and who are preferential
nose breathers [3, 4].
Eur Respir Mon, 2005, 31, 166–194. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
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PAEDIATRIC LUNG FUNCTION TESTS
Developmental changes
Developmental changes that influence the assessment of lung function in infants
include: 1) the compliance of the chest wall; 2) dynamic elevation of functional residual
capacity (FRC); and 3) the influence of the upper airways. During infancy, the highly
compliant chest wall results in minimal outward elastic recoil such that, during passive
expiration, the lungs recoil to a much lower volume in relation to total lung capacity
(TLC) than in older subjects. The potential difficulties imposed by the compliant chest
wall, including instability of FRC and a tendency for small airway closure during tidal
breathing, are partially compensated by dynamic elevation of the end expiratory level.
During the first months of life, infants modulate both expiratory time and flow to
maintain an adequate FRC. In addition to changes in respiratory rate, babies often use
laryngeal and post-inspiratory diaphragmatic activity to slow expiratory flow [4]. Unless
care is taken to limit recordings to periods of quiet sleep, the variability of end expiratory
level may impede assessment and interpretation both of lung volumes and of respiratory
mechanics and forced expiratory flows, which are highly volume dependent. Changes in
respiratory rate, expiratory time and the emptying time of the lung with growth may all
have significant effects on the interpretation of changes in various indices with growth [5–
7]. In infants, as in adults, nasal resistance representsy50% of total airway resistance. In
contrast to adults, however, infants are preferential nose breathers, with PFTs generally
being performed using a mask rather than mouthpiece and noseclips. Changes in intrathoracic airway resistance as a result of disease or therapeutic interventions may
therefore be masked, especially if there has been a recent upper respiratory infection. For
this reason it is usually necessary to postpone PFTs in infants for at least 3 weeks
following any respiratory infection.
Differences in measurement conditions
1) Sleep state, sedation and duration of testing procedure. Measurements in infants
are normally limited to periods of sleep so that the infant will tolerate manoeuvres such as
positioning of a face mask, brief airway occlusions and application of an inflatable jacket.
It is essential that a stable end-expiratory level is established before measurements
commence. To achieve this it is usually necessary for the child to be in quiet, rather than
rapid eye movement, sleep. Since the duration of such epochs are inversely proportional to
the post-conceptional (i.e. gestationalzpostnatal) age of the child [8], this presents a real
challenge when undertaking measurements in very young or immature infants.
Successful measurements using a full range of tests can usually be achieved during
natural sleep following a feed in all infants up to at least 1 month postnatal age (corrected
for any prematurity). Sedation is usually achieved using oral chloral hydrate in a dose of
50–100 mg?kg-1. With the exception of a small proportion of "high risk children" (such as
those with known or suspected upper airway obstruction [8]), chloral hydrate has been
shown to have an excellent safety record [9]. Nevertheless, its action can be
unpredictable, with time taken to fall asleep after administration ranging from 15 min
tow2 h, and duration of subsequent induced sleep can be equally variable. The relatively
short time that an infant spends asleep, with or without sedation, means that important
decisions have to be made regarding which respiratory function test(s) should be used.
2) Posture. During infancy, most PFTs are obtained in the supine position. This will
influence the position of the diaphragm, efficiency of respiratory musculature, FRC, lung
mechanics and the distribution of ventilation. Such changes must be taken into account
when interpreting results, particularly when reporting longitudinal changes.
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P.J.F.M. MERKUS ET AL.
3) Safety issues. Despite the excellent safety record, strict safety precautions must be
adhered to during all infant PFTs. Resuscitation equipment, including suction and
oxygen, needs to be available and two skilled operators, fully trained in basic life support,
one of whom has prime responsibility for monitoring the infant, must be present. Pulse
oximetry is recommended throughout the entire session.
4) Equipment requirements. During recent years, close collaboration between scientists
and manufacturers, guidelines published by a European Respiratory Society (ERS)/
American Thoracic Society (ATS) Task Force [1, 10–14], and technological advances have
made it possible to perform standardised infant PFT in an increasing number of centres.
5) Leaks and deadspace. The use of a mask rather than a mouthpiece may introduce
both physiological problems, due to the relatively large apparatus deadspace, and
technical problems, in that it is difficult to estimate the "effective" dead space of the mask
[15]. Leaks around the face mask, which occur frequently, can be difficult to identify.
Many centres use therapeutic putty to facilitate an airtight seal between face and mask.
Preschool children
Preschool children are too old to sedate for lung function testing, and may lack the
necessary coordination or concentration to perform some of the manoeuvres required for
lung function tests designed for older subjects. They also have a short attention span and
are easily distracted and thus need to be engaged and encouraged by the operator to
participate in the test. Thus, while measurement conditions for testing preschool children
are broadly similar to those required for older subjects, every effort should be made to
make the environment as child-friendly (and safe) as possible. This includes provision of
suitable furniture, games and wall coverings, as well as adaptation of the normal lung
function terminology.
To this end the most important acquisition for any "preschool set up" is personnel with
suitable temperaments, i.e. a love of children, infinite patience and stamina, and a good
sense of humour. Adaptability, meticulous attention to detail and a thorough
background in respiratory physiology are also essential requirements, since appropriate
criteria for acceptable tests in the preschool child may differ markedly from those
established for older subjects [16]. The criteria for a successful preschool session should
not only be that valid respiratory function results are obtained, but also that the children
and their parents want to return for subsequent visits. Because young children tire easily,
visits should be timed carefully to maximise success. The emotional and developmental
stage of the child are important determinants of success with preschool PFTs, and the
child’s past medical history may also be relevant. Those with extended hospital stays
during the neonatal period may display considerable antipathy towards any electronic
equipment or facial attachments. The need to gain the child’s confidence, provide
coaching in the various techniques and manoeuvres and accommodate a rest or games
between the different PFTs when necessary, means that plenty of time should be set aside
for preschool PFTs. The use of computer games and appropriate incentives to help the
child understand what is required can be helpful [16] while encouragement to sit quietly
during more prolonged periods of data collection (e.g. during multiple breath inert gas
washouts) can be provided by the opportunity to watch a favourite video.
Commercial equipment is available for most preschool tests, albeit not specifically
designed for this age group. The potential effects of using equipment that was developed
for older and larger individuals, particularly with respect to deadspace and resistance,
168
PAEDIATRIC LUNG FUNCTION TESTS
need to be considered. These comments are equally relevant when assessing young
school-age children.
Anthropometry and background details
Given the rapid growth during infancy and early childhood, accurate measurements of
height and weight using a calibrated stadiometer and scales are essential. For accurate
interpretations of the lung function tests it is also essential to record data on
environmental, genetic and socioeconomic factors likely to impact on lung growth
including: sex; ethnic group; family history of asthma and atopy; cigarette smoke
exposure, both pre- and post-natal; allergen exposure, including pets; relevant current
and past medical history and medication use.
Purpose of assessing lung function in infants and children
Lung function tests in the very young are rarely performed for diagnostic purposes,
but rather to monitor the nature and severity of respiratory disease or to assess the
response to treatment. As in adults, lung function measurements in older children can be
used:
. as a diagnostic aid, to help determine the nature of the lung function disorder;
. to quantify the magnitude of the lung function disorder;
. to assist in determining prognosis or peri-operative risk;
. to assess the effects of medical interventions or diagnostic tests (such as the effects
of bronchodilator and bronchoconstrictor stimuli);
. to evaluate innovative therapies aimed at improving prognosis, quality of life and
lung function;
. to study the natural course of respiratory disease;
. to study the growth and development of the lungs and airways and evaluate early
determinants of airway function.
Though there is little doubt about the value of infant and preschool PFTs in
epidemiological and clinical research studies, their potential influence on individual
clinical management is more debatable [17–19]. Within infants, the clinical utility of such
tests will be limited by the time consuming nature of these tests and the need for sedation
beyond the neonatal period, whereas the availability of "preschool tests" has been too
recent for any reliable assessment of the potential value of such assessments within
individuals. While highly reproducible measurements of lung function can be made in
infants and young children during the same test occasion, little is yet known about the
"between test repeatability".
During childhood, beyond the neonatal period, lung function disorders are usually of
an obstructive nature, generally being confined to the intrathoracic, intrapulmonary
airways. Hence, measurements of airway patency (maximal expiratory flow volume
(MEFV) curves, spirometry, resistance measurements) are of most relevance for these
patients. Restrictive lung diseases are less common in children than in adults, such that
measurements of static lung volumes, maximal inspiratory or expiratory pressures, lung
compliance and diffusion capacity are used less frequently. Such tests may however play
a role in children suffering from cardiac diseases, systemic vasculitis, immunological
disorders, neuromuscular and/or orthopaedic disease. Similarly, tests designed to assess
parenchymal lung disease (lung compliance, lung volumes and partitioned respiratory
mechanics) will be particularly pertinent in those delivered prematurely and/or suffering
from the respiratory distress syndrome or chronic lung disease of infancy. During the
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P.J.F.M. MERKUS ET AL.
past decade, development of noninvasive methods, such as analysis of exhaled nitric
oxide and breath condensates to assess the metabolic function of the lung, have begun to
play a role in diagnosing and monitoring airway inflammation, especially in allergic
asthma. The following section summarises the techniques currently available to study
respiratory physiology and noninvasive inflammatory markers in children of various
ages.
Methods of assessing pulmonary function in infants and young
children
Details of most of the commonly used infant PFTs have been summarised previously
[1, 3, 10–14, 20], whereas those relating to preschool tests have been emerging at an
increasing rate during the past few years. Some of these tests can be applied throughout
childhood, whereas others are specific to either a sleeping infant or cooperative preschool
child. It is generally important to undertake tests based on tidal breathing recordings
prior to any forced expiratory manoeuvres. It is also essential to consider what is feasible
in the time available: while this is largely dictated by the duration of sleep in infants, and
that of concentration and cooperation in preschoolers, it also dependent on the age and
health of the subject and the expertise of the operators.
Even in older children, measurements may take considerably more time than in adults.
The lung function technician should report the child’s degree of cooperation. In young
children, the chances of successful initial attempts at spirometry may increase if they have
been introduced to the laboratory, and have practised blowing on a peak flow meter at
home before their first testing session. Given the wide range of body size, several sizes of
mouthpiece, and seats adjustable in height, should be available. Bacterial filters are
mandatory. Unless otherwise specified, any medication should be discontinued for a
prescribed period prior to testing; generally this would be 8 h for short-acting
bronchodilators, 48 h for long acting bronchodilators and 72 h for any anti-histamines
prior to a histamine provocation test.
Forced expiratory manoeuvres
While international guidelines are well established for spirometric measurements in
adults [21], the extent to which these are appropriate for school-aged children remains
questionable [22]. As discussed below, although major adaptations in terms of both data
collection and analysis are required, similar information from "full" MEFV curves can
now be obtained from both infants [20] and preschool children [16, 23–27].
The principles of paediatric spirometry in children aged w6–7 yrs are the same as in
adults (see Chapter 1). When testing preschool children, specially adapted quality control
criteria are required to allow for the fact that:
. Repeatability criteria need to be adjusted for the smaller absolute flows and
volumes being measured.
. Young children may need more attempts when learning to produce an MEFV
curve than are necessary in older subjects.
. Lung emptying occurs much faster than in older subjects. In preschool children
expiration may be complete inv1 s, making the use of a forced expiratory volume
in one second (FEV1) an inappropriate outcome measure. Indeed, even when the
child does produce a valid FEV1, its value may approach forced vital capacity
(FVC).
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PAEDIATRIC LUNG FUNCTION TESTS
With appropriate training and encouragement, most children 3–6 yrs of age can
achieve acceptable spirometry results. A variety of blowing games involving straws,
bubbles and party whistles can facilitate this process, as can demonstrations from the
operator and the use of carefully selected computer incentives [16, 28]. Considerable
input is required from the investigator with respect to the targets that are set: too low and
the child will not make a maximum effort, too high and they will become discouraged
and stop trying [16].
Spirometry can be performed with the subject seated or standing, but posture should
be reported. Most preschool children tolerate noseclips, although this is not mandatory
for acceptable recordings [29]. For quality control reasons all loops should be saved for
reviewing after the test. Three technically satisfactory curves can usually be obtained with
persistence.
Potential guidelines for spirometric assessments in preschool children have been
published recently [16]. Recommended criteria for acceptance of MEFV curves for
school children [22], include:
. that the total duration of forced expiration can be much less than the 6 s
recommended for adults (e.g. at least 2 s in children 8–19 yrs) provided there is a
asymptotic approach of flow versus volume, or volume versus time;
. that the difference between the two best FEV1s or FVCs should be based on a
percentage (rather than absolute) difference of the highest value.
Preliminary reference values have also been published for preschool children, but have
yet to be evaluated outside the centre where they were created [26–27].
Peak expiratory flow measurements
Widespread use of home peak flow monitoring has been considered a convenient and
cheap way to assess the pulmonary condition of children with asthma in the home
environment. More recent studies, however, seriously question the value and validity of
such recordings. It appears that peak flow diaries are often made up [30], are unreliable
and, most importantly, there is poor correlation between peak expiratory flow (PEF) and
measures of peripheral airway function [31]. The contribution of home PEF recordings to
better asthma management is unclear [32, 33]. The inter-subject variability for a specific
age or height is huge, which implies that a "personal best" value is the only reliable
anchor point. Peak flow diaries may still be of some use in isolated cases [34].
Forced expiratory manoeuvres in infants
By substituting an externally applied pressure to the chest and abdomen instead of
voluntary effort to force expiration, it has been possible to obtain both partial and "full"
forced expiratory manoeuvres in sleeping infants. Measurements of maximal flow at
FRC (V9maxFRC) have been used to help characterise the normal growth and development of the lungs during infancy and the pulmonary abnormalities associated with acute
and chronic lung disorders during early childhood. Data derived from forced flows and
volumes over an extended volume range have been found to discriminate clearly between
those with and without respiratory disease, even in the absence of symptoms [19].
Tidal rapid thoracic compression technique
In infants, partial expiratory flow volume curves can be produced by a jacket around
the chest and abdomen, which is inflated at the end of a tidal inspiration to force
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P.J.F.M. MERKUS ET AL.
expiration. The resultant flow is recorded through a flowmeter attached to a face mask
(fig. 1). This technique is usually referred to as the "squeeze" or the tidal rapid thoracoabdominal compression (RTC) technique. V9maxFRC is the most commonly reported
parameter derived from this technique (fig. 2) and equates to forced flows at low lung
volumes (e.g. MEF25%FVC) in older children [5]. Guidelines regarding data collection
and analysis for the tidal RTC have been published [11], as have sex-specific collated
reference data [35].
V9maxFRC is thought to reflect primarily airway calibre upstream to the airway
segment subjected to flow limitation and therefore to provide a measure of
intrapulmonary airway function that is relatively uninfluenced by the resistance of the
upper airways. This makes it a useful measure of intrathoracic airway function in infants,
in whom nasal resistance comprises such a large proportion of total resistance. As in
older subjects, the shape of the loop, as well as the derived numerical values, contribute
to interpretation of results (fig. 2). The tidal RTC technique has been used widely in
clinical and epidemiological research studies [19].
Large-bore
tubing and valve
Pressure relief
valve
Flow meter
Compressed
air
50 L
tank
Inflatable jacket
Fig. 1. – Equipment for partial forced expiratory manoeuvres using the tidal rapid thoraco-abdominal compression technique in infants.
Flow mL·s-1
a)
b)
200
100
V'maxFRC
0
-100
40
20
0
Volume mL
-20
40
20
0
Volume mL
-20
Fig. 2. – Partial flow–volume loops from a) a healthy newborn (maximal flow at functional residual capacity
(V9maxFRC)=92 mL?s-1) and b) a newborn with evidence of airway obstruction (V9maxFRC=40 mL?s-1).
172
PAEDIATRIC LUNG FUNCTION TESTS
Raised volume rapid thoraco-abdominal compression
The tidal RTC technique has been modified so that forced expiratory flows and
volumes can be measured over an extended volume range, in what has become known as
the "raised volume RTC or RVRTC technique". Similar to spirometric assessments in
older subjects, the raised volume technique allows the infant to inhale towards total lung
capacity before rapid inflation of the jacket initiates forced expiration from this elevated
lung volume, the manoeuvre ending when the infant reaches residual volume (RV).
The airway pressure used to augment inspiration is most commonly 30 cmH2O
(2.94 kPa). Jacket inflation must be maintained until lung emptying is complete. This
procedure is repeated until at least two technically satisfactory and repeatable
manoeuvres have been recorded [20]. The relationship between a partial forced
expiratory manoeuvre initiated from end tidal inspiration and that recorded from the
same infant after raising lung volume with an inflation pressure of 30 cmH2O is shown in
figure 3.
Parameters commonly reported from the RVRTC include the "forced vital capacity
from the applied inflation pressure", e.g. FVC30, FEV after 0.4 or 0.5 s and maximum
expiratory flow at 25% of forced vital capacity (MEF25). Calculations of FEV1 are rarely
feasible in young infants, except in the presence of marked airway obstruction, due to the
very rapid lung emptying and short forced expiratory time (FET) during early life [5, 6].
There is a marked age dependency of FEVt/FVC ratios during infancy and early
childhood and such ratios are poorly discriminative of changes in airway function due to
disease [36, 37].
Raised lung volume flow-volume manoeuvres are more difficult to perform than
partial flow-volume manoeuvres. Extensively trained personnel are needed to ensure
accurate results. Potential advantages of the RVRTC technique include the fact that
forced flows and volumes are measured from a reproducible, rather than a potentially
variable, lung volume; flows can be assessed over an extended volume range from near
TLC to RV; flow limitation should be easier to achieve; and longitudinal assessments of
similar parameters are possible from infancy to adulthood.
900
Flow mL·s-1
600
300
F=0
0
-300
150
100
50
Volume mL
0
-50
Fig. 3. – Overlay of forced expiratory flow–volume curves from the same infant obtained at the end of tidal
inspiration (inner curve; ?????) and after lung inflation to 30 cmH2O (outer curve; ––).
173
P.J.F.M. MERKUS ET AL.
Several studies have indicated that RVRTC may be more discriminative than tidal
RTC for distinguishing the effects of respiratory disease on airway function [38, 39]. This
technique has the potential to quantify the degree of airway obstruction [38, 40], monitor
changes in airway mechanics over time [36, 37] and evaluate bronchial responsiveness
[40–42]. Measures of FEVt are more reproducible than forced flows and have been found
to be discriminative [36, 37, 43]. Forced flows may, however, be sensitive in wheezy
infants during both baseline measurements and assessments of bronchial responsiveness
[40, 41, 44].
Resistance and compliance measurements
Measurements of resistance reflect airway function during tidal breathing and are thus
suitable for subjects who cannot actively participate in lung function tests. This section
summarises special issues to consider when undertaking these measurements in infants
and children, as well as providing an overview of techniques that have been developed
specifically for use in younger subjects.
Plethysmographic assessments of airway resistance
Whole body plethysmography has been successfully adapted for measurements in both
sleeping infants [3, 14] and awake preschool children [45]. Airway resistance (Raw) has
been used to study normal airway growth and development in relation to lung volume
during the first year of life and has demonstrated the presence of "tracking" of airway
function during this period [46, 47]. These measurements have also been used to
discriminate between healthy infants and those with respiratory disease or prior wheezing
[46, 48, 49]. Plethysmographic measurements of specific airway resistance (sRaw: i.e.
Raw6FRC) are becoming an increasingly popular method of assessing both baseline
airway function and bronchial responsiveness in preschool children. Values of specific
resistance remain relatively independent of changes in body size. This should facilitate
attempts to distinguish changes in airway function due to disease from those resulting
from growth and development. sRaw has been used as an outcome measure in healthy
young children [45, 50] and those with cystic fibrosis [51, 52], as well as being used to
assess bronchial responsiveness or document the effect of anti-asthmatic therapies in
preschool children [53–55].
The principles of plethysmographic assessments of airway resistance measurements are
identical in infants and preschool children to those in older subjects [3, 14]. Most sleeping
infants will readily tolerate an airway occlusion lasting 6–10 s, and will make respiratory
efforts against the shutter, thereby generating the necessary pressure and volume changes
to calculate FRCpleth.
Most commercially available "adult" plethysmographs now rely on some form of
electronic compensation, rather than a heated re-breathing system and/or the need
for panting manoeuvres to minimise the influence of changes in humidity and
temperature of the respired gas during these measurements. This method has
been successfully adapted to measure sRaw in preschool children [45], but has yet to
be validated in infants, in whom published data has largely been obtained while
breathing gas under body temperature, saturated (BTPS) conditions [3, 46, 48].
In preschool children, who are unlikely to tolerate the relatively prolonged airway
occlusion required for measuring FRCpleth, sRaw is obtained simply by measuring
changes in air flow relative to changes in plethysmographic volume during spontaneous breathing, i.e. as the slope of the specific resistance loop without simultaneous
174
PAEDIATRIC LUNG FUNCTION TESTS
measurement of FRC.
sRaw~(DV pleth=DFlow)|(Pamb{PH2 O)
ð1Þ
Where Pamb is ambient pressure and PH2O is water vapour pressure. One of the
shortcomings of plethysmographic measures of airway resistance is that there is no
consensus regarding which parameters should be reported. Since resistance changes
throughout the breathing cycle, there is no single value that can be considered truly
representative. Most commercially available systems have several ways of calculating
resistance and specific resistance including the pressure–flow relationship:
. between points of maximum pressure swing or maximum flows;
. throughout the entire breath (calculated by regression of DVpleth/DFlow); and
. at a fixed flow during initial inspiration and/or expiration.
Possibly, the effective Raw (Reff) is the only objective and standardised way of
calculating Raw because it is representative for the entire breathing cycle and found by
dividing two integrals. It is possible to separate into Refftotal, Reffin and Reffex.
2
Reff ~integral (Pa:V 0 :dt)=integral((V 0 ) :dt)~integral(PA:dV )=integral(V 0 :dV ) ð2Þ
The interpretation is, however, difficult. The major advantages of plethysmographic
Raw are that it represents a direct reflection of airway calibre during tidal breathing
and that a similar method can be used at all ages. Infant plethysmography, while
providing valuable data in specialised centres, remains limited by the lack of any
validated method of obtaining reliable results without reliance on a heated rebreathing bag, and the potential dominance of the upper airway in these nose breathing
subjects. Preschool plethysmography has the advantage that the necessary equipment
is available in most secondary hospitals, thereby simplifying a wide dissemination of
this method, although its size precludes use in field studies and most primary care
facilities. Improvements in commercially available software are required to facilitate
both standardised data collection and quality control in this age group.
Passive respiratory mechanics
Measurements of passive respiratory mechanics (compliance, resistance, and
expiratory time constant) are possible if a state of relaxation can be induced in the
respiratory system. The vagally mediated Hering Breuer inflation reflex is active within
the tidal volume range throughout the first year of life. The "occlusion technique" for
measuring passive respiratory mechanics is based on the ability to invoke this reflex by
performing brief intermittent post-inspiratory airway occlusions during spontaneous
tidal breathing. This leads to inhibition of inspiration, and prolongation of expiratory
time (fig. 4). Provided there is no respiratory muscle activity and that there is rapid
equilibration of pressures across the respiratory system during the occlusion, alveolar
pressure can be measured at the airway opening. By relating this recoil pressure either to
the volume above the passively determined end expiratory volume at which the airway
occlusion was performed or to the airflow occurring on release of the occlusion, the
compliance and resistance of the respiratory system can be calculated [12]. A popular
adaptation of this technique is the single occlusion technique (SOT). When using this
technique, resistance, compliance and the passive expiratory time constant (trs) of the
respiratory system can be calculated from a single airway occlusion (fig. 5). Provided
expiration is passive and there is no braking of expiratory flow following release of the
airway occlusion, a plot of the flow-volume relationship can be used to calculate t (since
time constant=volume/flow). Compliance of the respiratory system (Crs) is calculated by
relating the volume above the passively determined lung volume at the moment of airway
175
P.J.F.M. MERKUS ET AL.
Elastic recoil
pressure
V
Pao
Inspiration
Time s
Fig. 4. – Assessment of passive mechanics using the occlusion technique, which is based on the ability to invoke
Hering-Breuer inflation reflex. Airway occlusion at end tidal inspiration induces a respiratory pause (lengthening
of expiratory time), during which the recoil pressure of the respiratory system can be measured at the airway
opening. V: volume; Pao: airway opening pressure.
Pressure
P1
Flow
Fext
Vol%A
Time
Compliance = volume/pressure = Vext/P1
Resistance = pressure/flow = P1/Fext–Rapp
Vol%B
Volume
100%
V0 Vext
65%
5%0%
Fig. 5. – Assessment of passive respiratory mechanics using the single breath occlusion technique. Provided there
is passive expiration, the time constant of the respiratory system can be measured by regression of volume/flow
over the linear portion of the flow volume plot (i.e. between Vol%A and Vol%B) during expiration (a).
Compliance is measured from the ratio of volume in the lung above the passively determined end expiratory
level at time of occlusion (the "extrapolated volume" or Vext), divided by the elastic recoil pressure at that
volume (P1 – b), whereas resistance is measured either from the pressure/flow relationship as shown or as:
resistance=time constant/compliance (for further details see [10]). Fext: flow at Vext.
occlusion to the elastic recoil pressure measured during the occlusion. The respiratory
time constant represents the product of resistance and compliance. Respiratory
resistance (Rrs) can thus simply be derived as Crs/t. An occlusion time of at least
400 ms (maximum 1.5 s) in which to attain a pressure plateau lasting at least 100 ms has
been recommended [12].
The advantages of the SOT are that the equipment is relatively simple and cheap,
consisting of a flowmeter and shutter attached to a face mask, and that measurements
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PAEDIATRIC LUNG FUNCTION TESTS
can easily be made at the bedside. As with all infant PFTs, attainment of a stable
respiratory pattern and leak free seal around the mask are mandatory. Valid
measurements also depend on three vital assumptions, namely that:
. there is complete relaxation of the respiratory system not only during the occlusion,
but during the subsequent expiration;
. pressure measured at the facemask equilibrates rapidly with alveolar pressure; and
that;
. the lung can be treated as a single compartment, that can be described by a single
time constant.
These conditions can be achieved in most healthy infants during quiet sleep but may be
more difficult to satisfy in infants with severe airway disease. As with all measures of
resistance during infancy, results may be dominated by the upper airways, particularly if
there is any evidence of an upper respiratory tract infection. While significant changes in
Rrs have been reported among groups of infants with airway disease [56], the major role
for these measurements is to assess restrictive pulmonary changes in conditions, such as
respiratory distress syndrome, chronic lung disease, pulmonary hypoplasia, interstitial
lung disease and cardiac disease with pulmonary over perfusion [57–59].
Interrupter technique
Resistance of the respiratory system can be assessed in preschool children by using the
interrupter technique, which relies on short interruptions to airflow. The interrupter
technique was first reported in 1927 by von Neergaard and Wirz, who applied a sudden
brief occlusion to the airways (100 ms) during a normal respiratory cycle while recording
flow and mouth pressure (Pm). Based on the assumption that pressures equilibrate
rapidly throughout the respiratory system during periods of no airflow, such that Pm will
reflect alveolar pressure during the occlusion, the interrupter resistance (Rint) can be
calculated by dividing the change in Pm after the occlusion by the flow immediately prior
to the occlusion. Interest in the interrupter technique has been heightened during the past
decade, as its potential use as a clinical tool for measuring lung function in young
"noncollaborating" children has been appreciated [60].
Theoretically, when airflow at the mouth is suddenly interrupted, there will be a rapid
initial change in Pm (Pinit) followed by a slower change (Pdif) up to a plateau (Pel) (fig. 6).
Pinit is virtually instantaneous and reflects the pressure difference due to the airway
resistance at the time of interruption. During tidal breathing, Pinit, and thus Rint, will
include a component of lung tissue and chest wall resistance, not just airway resistance.
Pdif is due to the visco-elastic properties of the respiratory tissues and reflects stress
adaptation (relaxation or recovery) within the tissues of the lung and chest wall, plus any
gas redistribution (Pendelluft) between pulmonary units with different pressures at the
time of interruption. The final plateau usually represents the pressure due to the elastic
recoil of the respiratory system and may take several seconds to be reached, especially in
the presence of any airway obstruction. In reality such a plateau is rarely observed during
the interrupter technique due to the brevity of the shutter closure. The total time of
interruption should be v100 ms, to prevent breathing against the occlusion [60]. The
major advantage of the interrupter technique is its portability and the simplicity of data
collection, which makes it suitable for use in field work.
Reference values for interrupter resistance in preschool children have been published
[61–63], but methods and equipment used in different laboratories are not standardised.
The definition of a clinically significant decrease in Rint in response to a bronchodilator
[62, 64], the role of Rint in challenge tests and the usefulness of the interrupter technique
in comparison with other techniques remain to be determined.
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2.0
Pmo
1.5
Pdif
1.0
0.5
0.0
0.0
Pel
Pinit
0.5
1.0
Time s
1.5
2.0
Fig. 6. – Schematic description of the pressure–time curve showing mouth pressure (Pmo) changes after a sudden
interruption of airflow at mid expiration. Pinit: rapid initial change in Pmo; Pdif: secondary slower change in
Pmo; Pel: final plateau representing the pressure due to the elastic recoil of the respiratory system.
Forced oscillation technique
The forced oscillation technique (FOT), in which impedance of the respiratory system
is measured by superimposing small amplitude pressure oscillations on the respiratory
system and measuring the resultant oscillatory flow is another technique that has been
successfully adapted for use in infants and preschool children, based on tidal breathing.
A full description of data collection and analysis, together with guidelines for the
application and interpretation of FOT have been provided [65], and are discussed further
in Chapter 5 of this Monograph. The FOT can be used to define respiratory system
impedance (Zrs), if transrespiratory pressure is measured. Important limitations of the
technique include the effects of upper airway compliance. This is particularly important
in small children who have relatively low upper airway wall impedance relative to Zrs,
such that the latter may be underestimated.
Although some reference data have been reported [65], the wide degree of variability
between healthy children limits the extent to which this technique can be used to assess
either the presence or severity of airway disease within individuals. Nevertheless, in
asthmatic children, parameters derived from the FOT may provide a more reliable
evaluation of bronchial obstruction and its reversibility than Rint [66]. In addition, the
FOT facilitates noninvasive assessments of the bronchomotor response to deep
inhalation, as a reflection of the degree and nature of airway obstruction [67, 68].
These characteristics, together with the minimal requirement for the subject’s
cooperation, make FOT a suitable paediatric lung function test for epidemiological
and field studies.
Assessment of lung volumes and ventilation
Tidal breathing parameters
Tidal breathing parameters have been used in both clinical and research settings to
determine tidal volume, breathing frequency and minute ventilation, to investigate the
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PAEDIATRIC LUNG FUNCTION TESTS
control of breathing, to trigger equipment and as an indirect measure of airway
mechanics. Such measurements and their interpretation are in fact highly complex [10,
69] and it is therefore important to appreciate the numerous factors that may influence
these recordings.
Patterns of tidal flow-volume loops can potentially yield important information about
the likely site of obstruction (fig. 7). Attempts to quantify such patterns have resulted in
numerical descriptors of the tidal flow pattern, such as the time to peak tidal expiratory
flow as a ratio of total expiratory time (tPTEF/tE), which may be reduced in the presence
of airway obstruction. tPTEF/tE has been shown to be a valuable outcome measure in
epidemiological studies designed to investigate early determinants of airway function [47,
70–73]. However, this measurement is only distantly related to airway function and, as
with most tidal breathing parameters, conveys mixed information on the interaction
between control of breathing and airway mechanics [69, 74, 75].
The greatest advantage of undertaking these measurements is their noninvasive nature
but it is difficult to record baseline values of tidal breathing when using any system that
requires facial attachments. While attempts have been made to use body surface
measurements, results have often been disappointing [76, 77]. In health, the pattern of
tidal breathing is highly variable [75], and while assessments have been performed in
awake newborn infants, repeatable measures normally require that the infants are in
quiet sleep. In awake preschool children, such problems are even more marked. The
clinical usefulness of tidal breathing measurements is limited by the marked within and
between subject variability of breathing pattern. Within epidemiological studies, the
discriminative ability of indices such as tPTEF/tE decreases with increasing age [3, 46, 73,
77].
Measurement of static lung volumes
Introduction. Measurement of static lung volumes may be essential for accurate
interpretation of volume-dependent pulmonary mechanics, such as lung compliance,
resistance or forced expiratory flows, as well as for defining normal lung growth. The most
common abnormality of lung volume during infancy and early childhood is associated
with airway obstruction, wherein both hyperinflation and/or gas trapping result in
elevated values of FRC.
Lung volume measurements in this age group have primarily been undertaken during
tidal breathing, i.e. FRC using either plethysmographic or gas dilution/washout
techniques. With the raised volume technique it is now possible to calculate quasi-values
of "TLC", "expiratory reserve volume" and "RV" [38].
Plethysmographic assessments of functional residual capacity. Assessment of static
lung volumes in children aged 3–5 yrs are generally limited to those that can be obtained
using one of the gas dilution or washout techniques, as they will rarely tolerate breathing
against the shutter.
Plethysmographic assessments of FRC are, however, well established in children aged
0–2 yrs [14, 38] and have been used in both clinical [40, 48] and epidemiological research
[46].
Functional residual capacity using gas dilution or washout techniques. Apart from the
differing measurement conditions discussed above and the need to miniaturise the
equipment, methods of assessing FRC by gas dilution or washout are much the same in
infants as in older subjects. Details of equipment specifications and techniques for infants
and preschool children have been published [13, 23].
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P.J.F.M. MERKUS ET AL.
a)
Insp
Exp
Volume
VT
tI
tE
ttot
Time
b)
Flow
PTIF
tPTIF
PTEF
tPTEF
c)
Time
VPTEF
PTEF
TEF50
Flow
Exp
Insp
TIF50
PTIF
0.75
0.5
VT
0.25
0.1
Volume
Fig. 7. – Graphical presentation of the relationship between a) tidal volume and time; b) tidal flow and time;
and c) tidal flow and tidal volume. Insp: inspiration; Exp: expiration; VT: tidal volume; tI: inspiratory time; tE:
expiratory time; ttot: total time of one breathing cycle; PTIF: peak tidal inspiratory flow; PTEF: peak tidal
expiratory flow; tPTIF: time to peak tidal inspiratory flow; tPTEF: time to peak tidal expiratory flow; VPTEF:
volume to peak tidal expiratory flow; TEF50: tidal expiratory flow at 50% of tidal volume; TIF50: tidal
inspiratory flow at 50% of tidal volume.
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PAEDIATRIC LUNG FUNCTION TESTS
During recent years, there has been increasing emphasis on the use of washout
techniques in infants and young children, using either the bias flow nitrogen washout
technique, which is based on a mixing chamber technique [13] or the multiple breath gas
washout (MBW) technique. The latter measures breath-to-breath changes in gas
concentration during the washout of an inert gas and provides information on both lung
volumes and ventilation distribution (see below) [78, 79]. Equipment designed for older
subjects can often be adapted for use in children, provided care is taken to minimise
deadspace of the circuitry.
Multiple-breath inert gas washout. The MBW test is performed during tidal breathing. The original test was the N2 MBW test, using 100% O2 for the washout, where
washout of N2 is monitored after inspiring 100% O2. This is a valuable and simple
technique for use in preschool and older children. The use of 100% O2 may, however,
alter tidal breathing patterns in young infants and is therefore less suitable in this
age group, particularly if measures of ventilation inhomogeneity are required. A
nonresident inert marker gas, such as Helium (He) or Sulfurhexafluoride (SF6), may be
used instead. The wash-in phase consists of the subject breathing a gas mixture containing
the tracer gas through a facemask until equilibration is achieved throughout the lungs.
The gas supply is then disconnected during an expiration so that "washout" can
commence with the subject breathing room air until the end-tidal tracer gas concentration is v1/40th of the starting concentration. The gas concentration and flows are
measured continuously at the airway opening. From such a washout, both the FRC and
indices of ventilation inhomogeneity (see below) can be calculated [78, 79]. Since the
MBW technique simply requires the child to breathe tidally through a facemask or
mouthpiece attached to a flow meter and gas analyser, it is eminently suitable for subjects
of all ages, from birth to old age.
With their rapid respiratory rate and higher ratio of tidal volume/FRC, wash-in and
washouts are generally much faster in this age group, both phases of the technique being
completed within 1–3 min in healthy subjects (the faster times being observed in infants)
and within 5 min in those with airway disease.
Disadvantages of the MBW or gas dilution/washout techniques include the fact that
they only measure the readily ventilated gas volume, and may require prolonged washout
in those with severe disease, especially in older subjects. However, these techniques are
suitable for bedside measurements, can be undertaken at all ages and can provide
simultaneous assessments of gas mixing indices.
Gas mixing efficiency
The use of the MBW to assess gas mixing efficiency or ventilation inhomogeneity
has only been used intermittently in infants and young children [80, 81]. Indices of
ventilation inhomogeneity have been shown to be increased in patients of all ages with
respiratory disease. MBW seems a more sensitive method of detecting early changes
in lung function among infants and children with cystic fibrosis than conventional
techniques [51, 79, 82]. A further advantage of this method is that gas mixing efficiency
remains remarkably stable throughout life, facilitating improved discrimination
between health and disease. There are as yet no guidelines for standardised
measurements of ventilation inhomogeneity in infants, but the use of the technique in
preschool [51] and school-aged [83, 84] children has recently been described in some
detail.
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P.J.F.M. MERKUS ET AL.
Diffusion technique
The passage of gases across the blood-gas barrier in the alveoli can be described by the
diffusion capacity [85]. Recommendations for standardised measurements in adults and
older children have been published [86], but recent guidelines for younger children are
lacking. Paediatric diffusion capacity measurements have been recently reviewed [87].
For children who cannot perform the single breath manoeuvre or who have a lung
volume v1.5 L, a rebreathing method can be used where the decay of CO is monitored
continuously in a closed system while the child breathes quietly (fig. 8, [88]).
Indications for assessment of diffusion capacity in children include monitoring during
and after chemotherapy or irradiation, diagnosis and monitoring of interstitial lung
disease, and monitoring for pulmonary bleeding disorders. One important limitation of
diffusion measurement is unreliability in the case of severe airway obstruction due to
inadequate time for equilibration of the gases in the airways. At reduced lung volume,
diffusion per unit lung volume increases and this may lead to erroneous interpretation of
data in children with restriction of chest movements. In that case, the use of appropriate
reference values, obtained at the relevant lung volumes, is essential. Falsely high
diffusion may be found due to the presence of blood in the airways and alveoli, or in the
case of relative hyperperfusion.
Noninvasive monitoring of inflammatory markers
In the last decade new, rapid and simple techniques have been developed to assess
inflammatory markers in exhaled air, which are especially attractive for paediatric
pulmonology. The technique for measuring exhaled nitric oxide (eNO) in older children
has been well standardised [89].
Measurement of exhaled nitric oxide
The most accessible and easy marker of airway inflammation in allergic asthma is the
fractional concentration of Nitric Oxide (FENO) in a sample of exhaled gas. Practical
Fig. 8. – Rebreathing method in a closed system for measurements of diffusion capacity during quiet breathing.
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PAEDIATRIC LUNG FUNCTION TESTS
recommendations and suggestions for measuring FENO in children are available [89], the
choice of method depending on the age and cooperation of the child.
School age children
Single breath on-line measurement. The single breath on-line (SBOL) method is
considered the preferred method for cooperative subjects, and is generally applicable in
children agedw5 yrs. The child should be seated comfortably and encouraged to breathe
quietly fory5 min. The child inhales to near-TLC and immediately exhales at a constant
flow of 50 mL?s-1 until an NO plateau of at least 2 s can be identified during an exhalation
of at least 4 s. The exhalate is sampled continuously and fed into a chemiluminescence NO
analyser. The inspired gas should contain minimal NO (v5 ppb). The expiratory pressure
should be maintained between 5 and 20 cmH2O to close the velum. This is necessary to
avoid contamination with nasal gas, which has a high NO concentration. Repeated
exhalations (three which agree within 10% or two within 5%) should be performed with at
least 30 s intervals, and mean NO recorded [89]. A target expiratory flow of 50 mL?s-1 has
a good reproducibility and discriminatory power in children [90, 91]. The SBOL technique
may be difficult in preschool children who often have difficulties in maintaining flow or
pressure within the required limits [92–94]. Audiovisual aids to facilitate inhalation to
TLC and control of expiratory flow, together with the use of dynamic flow restrictors that
allow the child to exhale with a variable mouth pressure while maintaining a constant
expiratory flow, are helpful [94]. Dynamic flow restrictors are simple manual or
mechanical devices that vary their resistance depending on the forced expiratory pressure
and their use is highly recommended in children.
Off-line method with constant flow and dynamic flow restriction. When there is no
analyser available, exhaled gas samples can be obtained and analysed later. Such exhaled
air samples are stable for up to 9 h after collection, and can be transported to the lab. The
child blows through a mouthpiece into a NO-inert balloon. Nasal contamination is
prevented by exhaling through a resistance that generates an oral pressure of at least 5 cm
H2O to close the velum [92, 93]. The gas is collected in inert Mylar or Tedlar balloons.
Wearing a noseclip and breath holding potentially affect FENO and are not recommended
[95]. Flow standardisation improves the reproducibility of off-line methodology. With
identical flows, the off-line results are similar to those of on-line methods in school
children and adolescents [96]. A major improvement in off-line collection can be expected
from the incorporation of a dynamic flow restrictor in the collection system, which is
feasible in children as young as 4 yrs [97]. It is recommended to use flows of 50 mL?s-1 for
both off-line and on-line collection [89].
Alternative methods for pre-school children and infants
On-line measurements:. In children aged 0–2 yrs there is limited experience with on-line
tidal NO measurements and practical recommendations and suggestions for measuring
FENO in preschool children and infants are lacking. Measurements have been performed
during quiet, regular tidal breathing. Expiratory air collected via a facemask covering nose
and mouth (mixed air) is contaminated by ambient NO and NO from the nose. However,
such measurements have been shown to correlate with values obtained through oral
breathing. FENO may be measured on-line during spontaneous breathing in children aged
2–5 yrs while the expiratory flow is adjusted by changing the expiratory resistance.
Sources of variability include the characteristics of the breathing pattern, expiratory
flows, and the level of inflation. FENO measured during spontaneous breathing do not
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P.J.F.M. MERKUS ET AL.
equate SBOL measurements [98], and separate characterisation is required, including
definition of normal values in healthy children. The development of hand-held
miniaturised NO analysers will likely contribute to the use of this test.
Off-line measurements. Exhaled air can be collected during tidal breathing via a mouth
piece or a facemask which are connected to a non-rebreathing valve that allows
inspiration of NO free air to avoid contamination by ambient NO. Exhaled breath
samples are collected into an NO-inert bag fitted with the expiratory port once a stable
breathing pattern is present. Methodological issues need to be resolved with all these
approaches before these techniques can be recommended for routine use in infants and
preschool children
Breath condensate. Cooling of exhaled air causes condensation of water vapour. Breath
condensate can be analysed for the presence of inflammatory mediators and other
putative markers of inflammation, among which are hydrogen peroxide (H2O2),
leukotrienes (LT), prostanoids, thiobarbituric acid reactive products, and metabolites of
nitric oxide. The methodology for these measurements has not been standardised.
H2O2 is produced by inflammatory cells and pulmonary macrophages, and elevated
levels have been found in breath condensate of cigarette smokers [99], and in subjects with
respiratory disease [100–102]. In adult asthmatic patients, the H2O2 concentration in
breath condensate correlates with sputum eosinophilia, but not with hyper-responsiveness
[103]. Exhaled H2O2 seems to respond to antibiotic treatment in exacerbations of airway
infection in cystic fibrosis [102]. Reference ranges for exhaled H2O2 have been published for
school-aged children [104].
Condensate can be obtained by passing exhaled air through a cold tube, the material of
which should be appropriate for the retrieval of the substances under examination (e.g.
glass in the case of H2O2, teflon for cysteinyl LTs). Various types of glass tubes or vessels
have been used [100, 101, 105]. Cooling can be accomplished by countercurrent
circulating ice water in a double jacketed tube. Alternatively, condensate can be obtained
by blowing air through a glass vessel that is placed in liquid nitrogen, capturing water
vapor in the exhalate as ice on the walls of the vessel [105]. Frozen condensate can be
stored until analysis depending on the substance of interest. Its H2O2 content remains
stable for at least a month [104].
Airway responsiveness to bronchodilators and bronchoconstrictors
Assessment of the bronchodilator response is one of the most important clinical
investigations performed in older children and adults. The response to bronchoconstricting agents is less frequently examined and such measurements do not play a central
role in clinical management. They may be useful for confirming or excluding a diagnosis
of asthma in older subjects but the role of such assessments during the first five years of
life is less clear, due to the difficulty in both performing and interpreting these tests and,
until recently, the lack of available tests for use in preschool children. The effectiveness of
bronchodilators in wheezy infants remains controversial, reflecting the fact that, in
many infants who wheeze, the reduction in baseline airway function is not due to
reversible bronchoconstriction, but transient conditions associated with diminished
airway patency [41, 106]. In asthmatic children, the degree of bronchodilator response
appears to be age related, increasing from minimal or absent below 18 months of age, till
well established by 8 yrs of age [4]. It should be noted that many healthy preschool
children also demonstrate a significant response to bronchodilators. Responsiveness to
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PAEDIATRIC LUNG FUNCTION TESTS
bronchoconstrictors in infants may result from anatomically small airways or increased
smooth muscle tone, from relatively thicker airway walls, decreased chest wall recoil or
increased airway wall compliance. Together with the difficulties in estimating the dose of
agonist delivered to the lung, such factors make it virtually impossible to determine agerelated changes in bronchial reactivity during the first few years of life [107]. Bronchial
challenge tests are described in more detail in Chapter 8.
Methodological issues
Important issues to address in the assessment of airway responsiveness include the
technique used to assess the change in airway function, the agent used, the dosage and
delivery efficacy of aerosol, quantification of the airway response, and the potential
clinical implications of the test result. Routine tests for assessment in older children are
spirometry and plethysmography. In infants, most studies have used the partial or raised
volume RTC technique. Spirometry has been used in preschool children, but because of
its demanding nature this is usually limited to older children [108]. Commonly used tests
for preschool children include the interrupter technique, forced oscillation or
plethysmographic assessments of specific resistance, or more indirect methods [108–
111]. Technical and physiological problems can influence interpretation of airway
responsiveness in early life. Bronchoconstrictor-induced changes in FRC may result in
paradoxical changes in V9maxFRC and in underestimation of airway responsiveness,
whereas failure of rapid pressure equilibration during airway occlusions may invalidate
results from the interrupter or occlusion techniques.
Choice of provocative stimuli and mode of administration.
Histamine and methacholine are the most commonly used direct-acting agents,
whereas adenosine monophosphate and hypertonic saline have been used to provide a
more indirect and more physiological stimulus. Cold or dry air challenges have been used
in preschool children. Aerosol output and lung dose vary according to characteristics of
the nebuliser and the drug, the inspiratory volume, nasal versus oral inhalation and the
breathing and flow patterns. If the inspiratory flow is greater than nebuliser flow (as is
common above 6 months of age), the individual will entrain air that does not contain
aerosol, thereby lowering aerosol concentration [4]. However, because lung deposition
increases with age, a relatively constant dose is probably administered over the paediatric
age range. The dose of nebulised salbutamol/albuterol has commonly been 2.5 mg and,
when using pressurised metered dose inhalers with a spacer, the dose varies between 400
and 800 mg.
Evaluation of response
There is no consensus on how to evaluate bronchodilator responsiveness in an infant
or young child. Different approaches include: 1) comparison of the best pre and post
values; 2) comparison of the pre and post mean values from several replicates; and 3)
definition of a specific pre-post change for the individual based upon the variability of the
measurement in the population. The assessment of the within-subject variability between
occasions in the absence of any intervention is needed to correctly interpret such tests
[112]. When using forced expiratory flows to assess airway responsiveness in infants, the
provocative concentration of the agonist to produce a 30% reduction (PC30)
from baseline in V9maxFRC or MEF25 (FEF75) has been used. This decrease exceeds
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P.J.F.M. MERKUS ET AL.
the intra-subject within test variability of the measurement and is frequently associated
with a change in the flow volume curve from convex to concave [42].
Interpretation of lung function results in infants and children
Within an individual infant or child, the clinical usefulness of any lung function test
will always be enhanced if serial measures can be undertaken rather than a single
assessment, and if the choice of test is based on the question to be answered, clinical
reasoning and a knowledge of the suspected underlying pathophysiology, rather than
simply on the equipment that happens to be available in any given centre. Given the
marked influence of factors such as preterm delivery, intrauterine growth retardation,
sex, ethnic group and maternal smoking during pregnancy, it is particularly important to
take a careful history from the parents when performing such tests during early life.
Reference equations
Reference equations are essential to express pulmonary function in relation to that
which would be expected for healthy children of similar age, sex, body size and ethnic
group; to characterise and monitor disease severity; to expand knowledge regarding
growth and development; and to study mechanisms of normal and abnormal function
and the natural history of disease. The addition of age as an independent variable may be
particularly important to optimise reference equations during the pubertal growth spurt
[113–115], whereas separate reference equations, based on arm span and or ulna length,
are required for children in whom height cannot be measured accurately (e.g. those with
progressive scoliosis and/or neuromuscular disease) [116–119].
As a result of secular changes in both the age at which puberty commences [120] and
final height attained [121], regular updating of reference equations is required. Additional
trends in predicted values can also occur due to alterations in equipment and protocol
[15]. The choice of reference equations directly influences the interpretation of paediatric
lung function data and this can have a significant impact on patient care and research
[113, 114, 122].
Most lung function data are normally distributed or can be transformed to a normal
distribution, such that 90% of "normal" values are found within the range -1.64 and
z1.64 sd (with 95% within -1.96 and z1.96 sd). When studying adults, values outside
these ranges are considered "unusual" or "abnormal". In paediatrics, lung function
variables of healthy subjects and those with respiratory symptoms and/or disease often
overlap to such an extent that a "normal" lung function parameter does not exclude
disease. Clearly abnormal lung function parameters will, however, often – but not by
definition – be associated with symptoms and disease.
Z scores, or standardised deviation (SDS) scores, are defined as: Z=(observed value –
predicted value)/RSD, where RSD is the residual standard deviation of the reference
population. Z scores are by definition normally distributed with a mean of 0, and RSD of
1. Hence, the Z score indicates how many sds an individual or group is below or above
predicted for any given parameter. Z scores indicate how likely a result is to occur within
a "normal population" and how far removed the result is from that predicted, having
taken the natural variability of that parameter into account; they are useful for tracking
changes in lung function with growth or treatment, and allow comparisons of various
lung function results from different techniques. Several recent publications have reported
Z-scores that can be used in infants and young children [26, 35, 61, 123]. It is particularly
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PAEDIATRIC LUNG FUNCTION TESTS
important to avoid "back extrapolation" of spirometric reference data collected in older
children and adults as these will generally underestimate predicted values in the under
fives, resulting in loss of sensitivity with which to detect changes due to disease. This is of
considerable practical importance, since most commercially available equipment will
automatically default to pre-selected (or factory loaded) prediction equations based on
adult data. When selecting reference data with which to interpret paediatric lung function
results it is essential to check how appropriate those data are [124], including whether:
. the same equipment, technique, and methods of analysis were used;
. a comparable and sufficiently large population was studied, with even distribution
of age/body size;
. raw data are available for inspection;
. appropriate statistical techniques were used. It is of practical importance to agree
on the choice of reference equations on a national level and to aim for regular
updated reference values for that specific population, since evaluation of medical
treatment and study results heavily depend on that choice.
Conclusions
During the past decade there has been remarkable progress in the field of infant and
preschool lung function testing, with measurements once thought impossible to obtain
below the age of 5–6 yrs now being performed regularly in children as young as 3 yrs.
Commercially available equipment and international recommendations are now
available for most routine infant lung function tests. Forced expiratory manoeuvres
can be performed over the full lung volume range throughout infancy and the preschool
years, while noninvasive assessments of gas mixing efficiency that are applicable from
birth to old age offer the possibility of detecting early changes in airway function in
children with respiratory disease. Similarly, the development of reliable and childfriendly techniques to assess inflammatory markers in exhaled air has markedly
improved the possibilities of monitoring airway inflammation in asthma in recent years,
and these techniques may also have applications in other respiratory diseases. The
potential prognostic value of such tests within individual subjects is as yet unknown,
since it is only during the last few years that more routine assessments of lung function
have been possible throughout the preschool years. Longitudinal studies are currently
being undertaken in infants and preschool children with a range of respiratory diseases.
Interpretation of such data will require similar longitudinal measurements to be
performed among healthy children from birth to school age.
Lung function testing in infants and young children will, nevertheless, always present a
challenge and it is therefore essential that the purpose of any test is clearly defined at the
outset and that dedicated operators with the necessary expertise and patience are
available to undertake and interpret such measurements. Despite its vital role in clinical
and epidemiological research, given the need for sedation, the specialised equipment and
the difficulty in repeating measurements at frequent enough intervals, it is unlikely that
lung function testing will ever gain a routine place in the clinical assessment of infants
with respiratory disease. By contrast, provided guidelines and standardised protocols can
be developed that incorporate appropriate quality control criteria for young children,
together with reliable reference data and information regarding the relative sensitivity
and specificity of the various tests in differentiating children with and without respiratory
disease, lung function tests in preschool children could soon assume a similar role to
those used in their school age counterparts (in table 1 the feasibility of the various tests
are summarised). To achieve this goal, continuing international collaboration will be
187
P.J.F.M. MERKUS ET AL.
Table 1. – Illustrates the feasibility of various techniques according to (developmental) age. Age range listed is
rough indication
Technique
Tidal breathing
Respiratory mechanics (Rrs,
compliance, time constants)
Partial and raised volume RTC
technique (MEFV curves)
Spirometry (MEFV curves)
PEF
Raw
sRaw
Rint
FRC (plethysmography)
FRC (helium)
Forced oscillation
TLC and RV
Multiple breath gas washout: gas
mixing efficiency
Exhaled NO: single breath on-line
Off-line method with constant flow
and dynamic flow restriction
Breath condensate
Diffusion technique
Comments
Preterm infants
children 0–2 yrs
Children
2–3 yrs
Children
3–6 yrs
Children
6–18 yrs
X
X
X
X
X
X
X
X
X
(X)
X
X
X
(X)
X
X
X
X (composite)
X
Spontaneous sleep
or sedation
needed
X
(X)
X
X
(X)
(X)
X
X
(X)
X
X
X
X
X
X
X
X
X
X
X
X (5–6 yrs)
X (4–6 yrs)
X
X
Passive/minimal
cooperation, careful
timing needed
X
Some active
cooperation
possible
X
X
Full cooperation
possible
Rrs: respiratory resistance; RTC: rapid thoraco-abdominal compression; MEFV: maximal expiratory flow volume;
PEF: peak expiratory flow; Raw: airway resistance; sRaw: specific airway resistance; Rint: interrupter resistance;
FRC: functional residual capacity; TLC: total lung capacity; RV: residual volume; Brackets: measurements may
be feasible but validity not firmly established and/or likely success rate relatively low.
required together with input from manufacturers to ensure that the available equipment
and software is optimised for this very important age group.
Summary
. Measurements of ventilatory function can be carried out in children of almost all
ages, except between 2–3 yrs of age, which remains a very difficult age group to
assess.
. The type of measurements that are feasible strongly depends on developmental age
of the child, always requiring considerably more time and effort than
measurements in adults
. Methodological guidelines exist for most measurements in infants and schoolchildren, and are being developed for preschool children
. It is strongly recommended to work according to published guidelines, and to
choose appropriate reference equations
. Reliable reference data need to be established for young children aged v7 yrs.
Prediction of values for such children should never be based on those extrapolated
from older subjects.
Keywords: Infants, preschool children, respiratory function tests.
188
PAEDIATRIC LUNG FUNCTION TESTS
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121. Freedman DS, Khan LK, Serdula MK, Srinivasan SR, Berenson GS. Secular trends in height
among children during 2 decades: The Bogalusa Heart Study. Arch Pediatr Adolesc Med 2000;
154: 155–161.
122. Subbarao P, Lebecque P, Corey M, Coates AL. Comparison of spirometric reference values.
Pediatr Pulmonol 2004; 37: 515–522.
123. Jones M, Castile R, Davis S, et al. Forced expiratory flows and volumes in infants. Am J Respir
Crit Care Med 2000; 161: 353–359.
124. Quanjer PH, Stocks J, Polgar G, Wise M, Karlberg J, Borsboom G. Compilation of reference
values for lung function measurements in children. Eur Respir J 1989; 2: Suppl. 4, 121S–261S.
194
CHAPTER 10
Respiratory mechanics in the intensive
care unit
G. Polese*, A. Serra#, A. Rossi*
*Respiratory Division, Azienda Ospedaliera, Ospedali Riuniti di Bergamo, Bergamo, and
Intensive Care Unit, Azienda Ospedaliera, Ospedale "Carlo Poma", Mantova, Italy.
#
Newborn
Correspondence: A. Rossi, Unità Operativa Pneumologia, Ospedali Riuniti, Largo Barozzi 1, I-24128
Bergamo, Italy.
Assessment of respiratory mechanics can play a central role in the management of
critically ill patients undergoing artificial ventilation because of acute respiratory failure
(ARF) [1]. This is a condition defined by a rapid deterioration in pulmonary gas
exchange that may be due either to alterations in the mechanical properties of the
respiratory system leading to ventilation–perfusion mismatching or shunt, or to
neuromuscular insufficiency causing alveolar hypoventilation.
Assessment of respiratory function and mechanics is of crucial importance:
. to understand the pathophysiology of the disease underlying ARF;
. to assess the status and progress of the disease;
. to provide guidelines for therapeutic measures (positive end-expiratory pressure,
bronchodilators, fluids);
. to improve patient–ventilator interaction;
. to prevent ventilator-related complications;
. to plan the discontinuation of mechanical ventilation.
Despite the great importance of monitoring lung mechanics in ventilator-dependent
patients, these measurements are not regularly performed. This is probably due to the
general prejudice that these measurements are difficult to obtain in the intensive care unit
(ICU).
This chapter will be focused on respiratory mechanics rather than on general
respiratory physiology reflecting the attitude and expertise of the authors.
The purpose of this chapter is to review briefly the most common methods and
techniques for measuring and monitoring respiratory mechanics at the bedside of the
patient in the ICU.
We will go through a three point analysis: 1) measurements during controlled
mechanical ventilation, i.e. in the relaxed, passive patients; 2) evaluation of respiratory
mechanics during assisted mechanical ventilation; 3) the issue of the patient’s evaluation
in the weaning process from mechanical ventilation.
Before discussing the monitoring techniques, it is necessary to provide a brief overview
of the mechanical properties of the respiratory system.
The act of breathing is performed against several impediments: elastic, resistive,
viscoelastic, plastoelastic, inertial and gravitational forces, compressibility of intrathoracic gas, and distortion of the chest wall from its relaxed configuration. Despite its
apparent complexity, the dynamics of breathing have been satisfactory represented, at
least for clinical purposes, by a single-compartment model consisting of a rigid tube and
a compliant balloon [2, 3].
Eur Respir Mon, 2005, 31, 195–206. Printed in UK - all rights reserved. Copyright ERS Journals Ltd 2005.
195
G. POLESE ET AL.
The equation of motion for inspiration during spontaneous breathing is the
following:
Pmus~Pel,rszPresz(Pin)zPEEPi
ð1Þ
Pel,rs~E rs|DV
ð2Þ
Pres~Rtot|V 0
ð3Þ
Pmus represents the pressure developed by the inspiratory muscles, Pel,rs and Pres are
the pressure dissipated respectively to overcome the opposing elastic and resistive
forces, Pin is the inertial force, PEEPi is the intrinsic positive end-expiratory pressure,
i.e. the elastic pressure that is present if inspiration does not begin from the relaxed
functional residual capacity, Ers is the elastance of the respiratory system, Rtot is the
total resistance of the respiratory system (plus the endotracheal tube, if present); DV
is the inspired volume and V9 the inspiratory flow.
Since Pin is normally negligible it can be removed from the equation:
Pmus~(E rs|DV )z(Rtot|V 0 )zPEEPi
ð4Þ
During assisted ventilation the equation changes to:
Pmus{Pvent~(E rs|DV )z(Rtot|V 0 )zPEEPi
ð5Þ
Whereas during controlled mechanical ventilation (CMV), when the patient is
relaxed, the Pmus is 0 and all the pressure is developed by the ventilator:
Pvent~(E rs|DV )z(Rtot|V 0 )zPEEPtot
ð6Þ
Where PEEPtot is the sum of PEEPi plus the external PEEP applied by the ventilator.
Controlled mechanical ventilation
Measurements of respiratory function during CMV can be performed using different
techniques. In the following it will be assumed that the patient is a relaxed, passive
patient, i.e. a patient without a significant spontaneous respiratory activity. Monitoring
during CMV can be performed off-line or on-line (see table 1).
Monitoring off-line:
Interrupter technique. The most widely used technique for assessing respiratory
mechanics in patients during CMV is the rapid airway occlusion technique. This
procedure, albeit introduced at the beginning of the last century, gained popularity in the
Table 1. – Monitoring during controlled mechanical ventilation (CMV)
Off-line
Special ventilator facilities (e.g. occlusion buttons, analogue output from flow and pressure transducers)
Physiological equipment (e.g. interrupter valve, transducers, connectors, operators)
Interfering manoeuvres (e.g. airway occlusion, changes in the ventilator settings)
On-line
Mathematical models (z measuring equipment!)
CMV, relaxed patient
Continuous breath-by-breath analysis
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RESPIRATORY MECHANICS IN THE ICU
last 10–15 yrs after a series of studies that have elucidated the theoretical aspects of the
technique as well as its physiological basis [4, 5]. This technique requires either ventilators
equipped with special software options or a specific "button" to control the inspiratory
and expiratory valves or additional equipment (i.e. pneumotachograph, pressure
transducer, occlusion valve) inserted in line to the ventilator circuit. When applied at
the end of expiration (end expiratory occlusion, EEO, fig. 1) it provides a measure of the
static intrinsic positive end expiratory pressure (PEEPi,st), also known as auto-PEEP.
In patients with PEEPi flow stops abruptly before the next mechanical inflation,
producing a characteristic "truncated" appearance on the expiratory flow curve. In this
condition, during CMV, dynamic intrinsic PEEP (PEEPi,dyn) can be calculated by
measuring the amount of pressure (airway pressure) that need to be developed by the
ventilator to reverse the flow from expiration to inspiration. This can be easily assessed
by recording simultaneously flow and airway pressure and averaging many breaths by
superimposing flow–pressure loops and looking at the point where pressure tracings
cross the zero flow (fig. 2) [6]. PEEPi,dyn represents the lowest regional PEEPi which has
to be counterbalanced by the positive pressure of the ventilator to start inspiratory flow
[6].
If the rapid airway occlusion is applied just before the end of the inspiration (end
inspiratory occlusion, fig. 3) it enables measurement of most respiratory mechanics
parameters. As it can be observed in figure 3 the rapid end-inspiratory occlusion is
a)
1
Occlusion
L·s-1
0.5
0
-0.5
-1
b)
50
cmH2O
40
30
20
11.7
10
0
0
5
10
Time s
15
20
Fig. 1. – Representative record with measurement of intrinsic positive end-expiratory pressure (PEEPi) by endexpiratory airway occlusion (EEO) in a mechanically ventilated patient during controlled ventilation with
constant inspiratory flow. Records of a) flow and b) pressure at the airway opening. Inspiration is upward. At
the end of the tidal expiration, the expiratory circuit of the ventilator is occluded and airway pressure becomes
positive, reflecting the end-expiratory elastic recoil of the respiratory system due to incomplete expiration. The
value of PEEPi is provided by the difference between the EEO airway pressure plateau and atmospheric
pressure. Visual detection of the plateau on airway pressure provides direct evidence of absence of leaks in the
circuits, respiratory muscle relaxation, and equilibration between alveolar and tracheal pressure. Modified with
permission from [6].
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G. POLESE ET AL.
45
b)
10
40
9
35
8
7
30
Paw cmH2O
Paw cmH2O
a)
25
20
15
10
PEEPi,dyn
3
2
1
5
0
6
5
4
-1
-0.5
0
0.5
Flow L·s-1
1
0
-0.1
1.5
-0.05
0
0.05
Flow L·s-1
0.1
Fig. 2. – a) Airway pressure (Paw) versus flow diagram in one mechanically ventilated patient. Twenty-two
consecutive ventilatory cycles are superimposed. Inspiration is rightward. It can be noted that the start of
inspiration precedes the zero flow (?????). b) The target portion (dashed box) is magnified in figure 2b. Dynamic
intrinsic positive end-expiratory pressure (PEEPi,dyn) is measured as the difference between the value of Paw at
zero flow and the end expiratory pressure. The small end expiratory pressure reflects the resistive pressure due
to the end expiratory flow and is included in the value of PEEPi,dyn. Modified with permission from [6].
characterised by an immediate drop in airway pressure from a peak (Ppeak) to a lower
value (P1), followed by a slow decay to an apparent plateau achieved after 3–5 s [7].
From these manoeuvres it is possible to compute static elastance (Est) and interrupter
a)
1
L·s-1
0.5
Occlusion
0
-0.5
-1
cmH2O
b)
Ppeak
15
P1
Pplat
10
5
0
0
2
4
Time s
6
8
10
Fig. 3. – Tracings of a) flow and b) airway pressure in a representative patient in whom the end-inspiratory
occlusion during constant flow inflation has been performed. After end-inspiratory occlusion, there was an
immediate drop from its maximum value (Ppeak) to a lower value (P1) and then by a slow decay to an apparent
plateau (Pplat), which is achieved w2 s in this particular patient, although 5 s are recommended to achieve real
static condition (see text for more explanation). Modified with permission from [7].
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RESPIRATORY MECHANICS IN THE ICU
resistance (Rint), as well as Rtot, according to the following equations:
Pplat{PEEPi,st
E st~
VT
Rint~
Ppeak{P1
V0
ð7Þ
ð8Þ
Ppeak{Pplat
ð9Þ
V0
Where VT is tidal volume and V9 the flow immediately preceding the occlusion.
Calculations should be based on at least three manoeuvres. From the values of Rtot
and Rint the resistance of the endotracheal tube need to be subtracted to obtain the
Rtot,rs and Rint,rs. The resistances of the endotracheal tubes are markedly curvilinear
and can be computed from the following equation:
Rtot~
R~K1 V 0 zK2 V 02
ð10Þ
Where K1 (laminar flow) and K2 (turbulent flow) are two constants and can be
directly measured "in vivo" or obtained from published data computed "in vitro" and
V9 is the flow.
Relaxed expiration. Measurement of respiratory mechanics during relaxed expiration
may be useful, being potentially less influenced than inspiration by settings on the
ventilator. When the patient is sedated or completely relaxed the expiration is mainly
governed by the mechanical properties of the respiratory system. If expiration continues
below the tidal end-expiratory volume, dynamic pulmonary hyperinflation exists and can
be computed from the difference in volume between the end of relaxed expiration and the
end-expiratory volume of the preceding breath.
Normal subjects and patients with increased elastance show a smooth decrease in
expiratory flow throughout expiration, whereas patients with airflow limitation exhibit a
curvilinear pattern, convex to the volume axis [8], often associated with a sudden drop of
flow, when there is dynamic hyperinflation.
Expiratory flow limitation can be confirmed by the lack of change in flow with
modification in pressure at the airway opening or in the added flow resistance or with
application of a negative expiratory pressure (NEP). First, external resistance may be
increased [9] or decreased [10], for example, by adding an expiratory resistance or
removing the expiratory circuit of the ventilator. Secondly, pressure at the airway
opening may be increased [11] by setting PEEP by the ventilator or decreased with the
negative expiratory pressure technique [12].
Monitoring on-line
Despite the great importance of monitoring lung mechanics in ventilator-dependent
patients, the measurements previously illustrated are not continuous [13]. The rapid
airway occlusion interferes with the ventilator settings, requires a valve or a specific
"button" on the ventilator, and thus is not suitable for continuous monitoring.
Continuous monitoring enables the early detection of changes in patient’s status, thus
allowing a rapid therapeutic response, as well as the evaluation of its effectiveness. This
requires noninvasive, breath-by-breath monitoring and the microprocessor-based
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G. POLESE ET AL.
ventilator have shown the potential for continuous monitoring of respiratory mechanics
in ventilator-treated patients [14, 15].
The monitoring systems currently used rely on estimation techniques that are not up to
date: isovolume method for calculating resistance [16], measurements of dynamic
compliance at points of zero flow [11], and measurement of dynamic PEEPi [17]. All these
techniques are based on the assumption of the first-order model of respiratory
mechanics, i.e. on a linear model, while resistance and elastance are known to be volume
and airflow dependent.
Tracking respiratory parameters in time is possible using these mathematical models of
breathing mechanics and recursive estimation techniques [15, 16]. A parsimonious model
is needed because the performance of recursive methods for real-time identification
sharply deteriorates with increasing model complexity [18]. This leads to the selection of
the first-order viscoelastic model for on-line monitoring of breathing mechanics.
However, the parameters are allowed to change during the breath to provide a better
description of the data, thus accounting for the non-linear behaviour of respiratory
mechanics during artificial ventilation. The algorithm used in the literature is the
recursive least square (RLS), which has been recently modified to account for the nonlinear behaviour of respiratory mechanics during artificial ventilation [19, 20]. Briefly, the
current authors have adopted an RLS algorithm, combined with the classical first-order
model of respiratory mechanics and the continuous measurement of airflow and airway
pressure, to quantify respiratory mechanics in real time [6]. The method constructs, from
the recursive parameter estimates during inspiration, a weighted mean and standard
deviation of dynamic resistance, elastance and PEEPi. The mean values are updated on a
cycle-by-cycle basis to allow real-time monitoring of these clinical indexes in ventilatordependent patients with acute respiratory failure of different origins, including chronic
obstructive pulmonary disease (COPD).
Recently, Volta et al. [21] applied a different method using the least square fitting with
the first order model keeping resistance and compliance constant over the whole
breathing cycle. They applied the method in patients with and without expiratory flow
limitation. They concluded that data weighted on inspiration were acceptable in both
patient populations.
We can conclude that the proposed technique, although it has been limited so far to
patients without any respiratory activity, during controlled mechanical ventilation, is a
simple and robust tool for clinical use in the routine of the ICU.
Assisted mechanical ventilation
Work of breathing
To achieve normal ventilation, work needs to be performed to overcome the elastic
and frictional resistances of the lungs and chest wall. Determinants of work of breathing
(WOB) in ventilator-dependent patients are shown in table 2.
During controlled mechanical ventilation the external WOB is performed by the
ventilator and can be easily computed from the area subtended by the inflation volume
and applied pressure (either airway or tracheal). Changes in WOB during CMV reflect
changes in the mechanical properties of the respiratory system (provided that the
ventilator setting is unchanged).
During assisted mechanical ventilation due to the activity of the inspiratory muscles it
is necessary to use the oesophageal balloon to compute the WOB.
During assist-control ventilation, as proposed by Marini and coworkers [22, 23] and
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RESPIRATORY MECHANICS IN THE ICU
Table 2. – Determinants of work of breathing in ventilator-dependent patients
Patient’s abnormal mechanics
Low compliance
High flow resistance
PEEPi
Diameter of the endotracheal tube
Ventilator circuit, valves and devices
Ventilatory pattern
VT and V9E
Inspiratory flow rate and waveform.
PEEPi: intrinsic positive end expiratory pressure; VT: tidal volume; V9E: minute
ventilation.
Ward et al. [24], it is possible to measure noninvasively the WOB done by the patient by
comparing the pressure–volume (or pressure–time) relationship with that during CMV.
This technique of measuring WOB assumes the unproved supposition that the
mechanical properties of the respiratory system remain essentially identical during
"passive" and "active" conditions. Difference in WOB may be used to compare the effects
of the ventilator settings, response to therapy, such as bronchodilators [25, 26].
Measurement of WOB may help in deciding the appropriate level or type of ventilator
assistance and avoid both excessive or insufficient support. However, the superiority of
the measurement of WOB, over more simple measurements (such as peak or plateau
airway pressure, PEEPi, etc.), has not been proved. Although used in research, this
procedure does not enjoy wide popularity in clinical practice because of its
"invasiveness", apparent complexity and the difficulty of its interpretation.
Pressure time product
A significant limitation of measurements of respiratory work is that it may
underestimate the energy expenditure of the respiratory muscles during isometric
contraction. To overcome this problem many authors have suggested the use of the
pressure time product for the respiratory muscles. This requires the placement of an
oesophageal balloon and is calculated as the time integral of the difference between
oesophageal pressure (Poes) measured during assisted ventilation and the recoil pressure
of the chest wall measured during passive mechanical ventilation with VT and flow
setting identical to the assisted breaths. This can be easily performed during assist/control
ventilation, but not during other ventilatory modalities, such as pressure support
ventilation, when there is a great variability in VT and flow. This can be overcome using
the method proposed by Jubran et al. [27]. The recoil pressure of the chest wall (Pel,cw)
can be computed by multiplying the chest wall elastance, measured during passive
ventilation, by the volume signal. Then the pressure time product is calculated as the time
integral of the difference between Poes and Pel,cw. Another obstacle is the possibility that
the rapid drop in Poes with the beginning of inspiration is not due to the activity of the
inspiratory muscles but instead to the cessation of the activity of the expiratory muscles.
This can lead to an overestimation of the pressure time product. This can be avoided by
also placing a gastric balloon and measuring the transdiaphragmatic pressure (Pdi), i.e.
the differential pressure between gastric and oesophageal pressure. The time integral of
Pdi measures only the respiratory effort of one inspiratory muscle, the diaphragm (i.e. the
major inspiratory muscle), but does not need any correction for expiratory muscles
activity or the chest wall elastic recoil [28, 29].
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G. POLESE ET AL.
Weaning from mechanical ventilation
Airway occlusion pressure
The decrease in airway pressure 0.1 s after initiating an inspiratory effort against an
occluded airway (P0.1) has commonly been used as an index of neuromuscular
ventilatory drive [30–32]. This measure is not a great task to be performed in the ICU and
requires only a pressure transducer, a two-way valve and a data acquisition system. This
measurement is also automatically performed by some modern ventilators without any
direct physician’s intervention. Furthermore, during assisted ventilation, in some
ventilators (with a trigger delay around 100 msec) P0.1 can be measured breath by breath
by reading the pressure developed by the patient during the inspiratory effort needed to
trigger the mechanical breath [33, 34].
The value of P0.1 depends not only on the inspiratory centre output and the intact
neural pathway to the inspiratory muscles, but also to the electro-mechanical coupling
and the efficiency of the muscles. If an elevated P0.1 is present it undoubtedly indicates an
increase in the activity of the respiratory centre, but a low value may be related not only
to a reduced centre activity, but also to an impairment in one of the factors mentioned
above, particularly in the pressure generating capacity of the inspiratory muscles due to
weakness or fatigue. Furthermore, if PEEPi is present, i.e. there is dynamic
hyperinflation, the P0.1 measured at the airway opening did not take in consideration
the extra-pressure required to overcome PEEPi. P0.1 is usually increased in patients with
acute respiratory failure.
Monitoring of P0.1 in mechanically ventilated patients may help to identify patients
with high probability of successful weaning (P0.1 ƒ4 cmH2O) among those with high risk
of weaning failure (P0.1 i6 cmH2O) [35, 36], or may help to adequately set the level of
ventilatory support [37, 38].
Maximal inspiratory pressure
The inability of the respiratory muscles to sustain spontaneous ventilation is the
primary indication for the main therapeutic modalities in the ICU, namely mechanical
ventilation [39]. The balance between the maximal inspiratory muscle strength and the
mechanical load is the major determinant of the ability to sustain indefinite alveolar
ventilation. Respiratory muscle performance is one of the major issues in deciding the
timing and pace with which mechanical ventilation can be discontinued [40, 41].
Measurement of inspiratory muscle strength can be assessed by measuring maximum
inspiratory airway pressure (PI,max) while the patients makes a maximum inspiratory
effort against an occluded airway [42]. This manoeuvre requires the patient’s cooperation
and coordination. This is not an easy task in ventilator-dependent, acutely ill patients. To
obtain more reproducible measurements in ventilator-dependent patients, Marini et al.
[43] suggested the use of a unidirectional valve to permit exhalation while inhalation is
blocked. This permits exhalation proximal to residual volume, where PI,max is expected
to be higher. The highest values of PI,max are usually reached after 15–20 s of occlusion
[43]. Despite this manoeuvre values of PI,max measured in critically ill patients are usually
underestimated and show a poor reproducibility [44].
Breathing pattern
Abnormalities in respiratory frequency (f) and VT are common in acutely ill patients
admitted to the ICU. VT, f and minute ventilation are easy to measure in intubated
202
RESPIRATORY MECHANICS IN THE ICU
patients, whereas they are more difficult to obtain in spontaneously breathing subjects
due to the low tolerance of a face mask or mouthpiece, and moreover during noninvasive
ventilation due to the leaks always present. All modern ventilators have a "monitoring
area" in which breathing pattern parameters are continuously displayed. However, to
ensure the accuracy and reliability of volume measurements it is preferable to have a
simple handheld spirometer.
Rapid shallow breathing is a frequent finding in critically ill patients and it has been
related to respiratory muscles fatigue by some [45, 46] but not all [47, 48] investigators.
An elevated frequency often is the earliest sign of impending respiratory distress, and the
degree of elevation is proportional to the severity of the underlying lung disease [49].
Breathing pattern is commonly monitored during the titration of ventilatory support. In
particular f and VT are used to identify the appropriate level of pressure support [50–54].
Conclusions
Respiratory function testing, either off- or on-line, due to its simplicity, its
noninvasiveness and clinical usefulness should be a common practice in the critically
ill patients. Its use may help to adjust ventilator settings, medical treatment and assist the
clinician in the weaning process.
Summary
Assessment of respiratory mechanics can play a central role in the management of
critically ill patients undergoing artificial ventilation because of acute respiratory
failure (ARF). This assessment is of crucial importance to understand the
pathophysiology of the disease underlying ARF and to improve the patient–ventilator
interaction and the medical treatment of the disease.
Despite the great importance of monitoring lung mechanics in ventilator-dependent
patients, these measurements are not regularly performed.
The purpose of this chapter is to review briefly the most common methods and
techniques for measuring and monitoring respiratory mechanics on-line and off-line at
the bedside of the patient in the intensive care unit (ICU) and to be persuasive about
the usefulness and the feasibility of monitoring respiratory mechanics in the congested
rooms of the ICU. The chapter includes a three point analysis: 1) measurements
during controlled mechanical ventilation, i.e. in the relaxed, passive patients; 2)
evaluation of respiratory mechanics during assisted mechanical ventilation; 3) the
issue of the patient’s evaluation in the weaning process from mechanical ventilation.
Keywords: Acute respiratory failure, intensive care unit, mechanical ventilation,
monitoring, respiratory mechanics.
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