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PREFACE
Over the years, the International Commission on Radiological Protection (ICRP),
referred to below as ‘the Commission’, has issued many reports providing advice on
radiological protection and safety in medicine. Its Publication 73 is a general overview of the area. These reports summarise the general principles of radiological protection and provide advice on the application of those principles to the various uses
of ionising radiation in medicine and biomedical research.
Most of those earlier reports are of a general nature. More recently, the Commission has focused on addressing some more specific situations where difficulties have
been observed. A series of such reports was initiated at the Commission’s meeting in
Oxford, United Kingdom, in September 1997. On the recommendation of ICRP
Committee 3, the Commission has established a number of Task Groups to produce
reports on topical issues in medical radiation protection. The Commission tries to
ensure that reports on such problem areas are written in a style which is accessible
to those who may be directly concerned with these issues in their daily work, and
that every effort is taken to ensure wide circulation of such reports.
Several such reports have already appeared in print as ICRP Publications 84, 85,
86, 87, 93, 94, 97, 98, and Supporting Guidance 2. The present report continues this
series of concise and focused documents, and several further such advisory reports
are being prepared.
ICRP Publication 87, published in 2000, dealt with ‘Managing Patient Dose in
Computed Tomography’. In that report, multi-detector CT was mentioned in passing in a couple of sentences, but the technique was developing and there was little
experience of dose management with MDCT. By 2005, the situation had changed entirely, MDCT was rapidly replacing ‘conventional’ CT equipment, and at the Commission’s meeting in Bern, Switzerland, in September 2005, a Task Group on dose
management in multi-detector computed tomography was launched. Its terms of reference were to consider the rapid increase of this new and faster technique and the
ensuing new clinical applications, and to discuss parameters specific to MDCT that
systematically increase or decrease patient dose. The Task Group was requested to
complete its report rapidly.
The membership of the Task Group was as follows:
M.M. Rehani (Chair)
H.D. Nagel
M. Kalra
C. McCollough
Corresponding members were:
L. Collins
W. Kalender
The membership of Committee 3 during the period of preparation of this report
was:
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C. Cousins (Chair)
Y. Li
S. Mattsson
H. Ringertz
E. Vañó (Secretary)
J.-M. Cosset
J. Liniecki (Vice-chair)
L.V. Pinillos-Ashton
M. Rosenstein
Y. Yonekura
I. Gusev
P. Ortiz Lopez
M.M. Rehani
C. Sharp
This report aims to serve the purposes described above. In order to be as useful as
possible for those purposes, its style differs in a few respects from the usual style of
the Commission’s publications in the Annals of the ICRP.
The report was approved for publication by the Commission at its meeting in
Essen, Germany, in March 2007.
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SUMMARY POINTS
Chapter 1. MDCT technology
Modern CT scanners employ multiple rows of detector arrays allowing rapid
scanning and wider scan coverage.
All new CT systems have multiple detectors with a single or dual x-ray source,
and a number of new dose reduction tools have become available commercially.
Chapter 2. Radiation dose in MDCT
There are a number of aspects specific to MDCT that systematically increase or
decrease patient dose compared to single-detector-row CT scanners (SDCT).
Initial reports indicated increased patient doses for MDCT relative to SDCT;
more recent reports show comparable or decreased patient doses.
If the user selects settings for MDCT identical to those used in SDCT, there can
be an increase in patient dose. Settings must be determined appropriate to a specific scanner model.
Chapter 3. What considerations and actions affect patient dose?
There is potential for dose reduction with MDCT systems, but the actual dose
reduction achieved depends upon how the system is used.
It is important that radiologists, cardiologists, medical physicists, and CT system
operators understand the relationship between patient dose and image quality and
are aware that, often, image quality in CT is higher than that needed for diagnostic confidence.
Images of the highest quality are not essential for all diagnostic tasks, but rather
the level of quality (e.g., low noise, medium or low dose) is dependent on the diagnostic task.
Objective measures such as image noise or contrast-to-noise ratio may not completely capture all of the features relevant to making a correct clinical diagnosis.
Thus, determining ‘optimal’ image quality can be a complex task, as both quantitative metrics (e.g., noise) and observer perceptions are involved.
There is increasing awareness of how adapting exposure factors can contribute to
the management of patient dose. However, the rate at which technology is changing requires continued attention to management of patient dose.
Automatic exposure control (AEC) systems enable scan protocols to be applied
using measures related to image quality. If the image quality is appropriately
specified by the user, and suited to the clinical task, then there is a reduction in
patient dose for all but the obese patient. In obese patients, the dose is increased
to improve the image quality.
AEC does not totally free the operator from selection of scan parameters, and
awareness of individual systems is important. While CT systems without AEC
require operator selection of mA or mAs, AEC systems require understanding
of concepts such as noise index, reference mAs and reference images in order
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for AEC to be operated effectively. Understanding of some parameters, e.g., the
standard deviation of image pixels or noise index, is not intuitive and introduces
the opportunity for error.
The selection of image quality parameters in AEC systems is not straightforward.
There is lack of consensus on how image quality is to be specified and there are
significant differences in the way different companies achieve exposure control.
Operator knowledge of the system is important.
Chapter 4. Dose management in clinical practice
Justification of CT use is a shared responsibility between requesting clinicians and
radiologists. It includes justification of the CT study for a given indication and
classification of the clinical indications into those requiring standard dose CT
and those requiring only low dose CT.
Scanning parameters should be based on study indication, patient size, and body
region being scanned so that patient dose can be managed based on these
parameters.
Guidelines (selection criteria for CT examinations) are necessary so that inappropriate studies can be avoided. In addition, alternative non-radiation imaging techniques should be considered, when appropriate.
Training of requesting physicians and CT staff can help in the management of
scan indications, protocols, and patient dose.
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GLOSSARY
The following definitions apply for the purposes of the present publication:
Angular (x,y) mA modulation
Angular (x,y) mA modulation addresses the variation in x-ray attenuation around
the patient by varying the mA as the x-ray tube rotates in order to equalise the
photon flux to the detector. Typically the operator chooses the initial mA or
mAs value, and the mA is modulated upward or downward from the initial value
with a period of one gantry rotation.
Artefact
The appearance on the CT image of details not present in the scanned area in the
patient. The common artefacts are due to partial volume effect and to beam hardening. Both effects usually result in streaking artefacts, which are observed in
regions of high contrast when there is a sharp discontinuity in object density, such
as at air-tissue, air-bone and metal-tissue boundaries. Streaking will also arise
from mechanical misalignment within the scanner and, in clinical practice, from
patient motion and the use of high-density contrast media.
Automatic exposure control (AEC)
A device which automatically determines and provides the exposure needed to
produce a preselected image quality by sampling the x-ray intensity at the image
receptor.
Beam collimation
See collimation.
Beam pitch
See pitch.
Beam hardening
The process of filtration of a polychromatic beam by the preferential absorption
of lower energy photons in tissue, with a subsequent increase in effective energy.
Collective dose
An expression for the aggregate radiation dose incurred by a population,
defined as the product of the number of individuals exposed to a source and
their average radiation dose. The collective dose is expressed in man-sieverts
(man-Sv) and is intended solely as an instrument in the optimisation of radiation protection.
Collimation
Geometrical limitation of the extent of the radiation beam in the z direction.
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Combined angular and longitudinal (x, y, z) mA modulation
The use of both angular (x, y) modulation and longitudinal (z) modulation to vary
the mA both during rotation and during longitudinal movement of the patient
through the x-ray beam (i.e., anterior/posterior versus lateral, and shoulders versus abdomen). The operator must still indicate the desired level of image quality.
This is the most comprehensive approach to CT dose reduction because the x-ray
dose is adjusted according to the patient attenuation in all three planes.
Computed tomography dose index (CTDI) and Computed tomography air kerma
index
See Appendix A.
Computed tomography number
See CT number.
Contrast enhancement
Administration of intravenous or intra-arterial contrast media (typically containing iodine) to increase the visibility of low-contrast structures owing to the
increased attenuation of vessels and organs/tissue containing contrast media.
Contrast-to-noise ratio
Contrast-to-noise ratio is the difference between CT numbers of any two regions
in the image divided by noise.
Note: This quantity needs to be introduced because attention only to the ‘contrast’ has often resulted in images of higher quality than needed for confident
diagnosis. Noise is also a measure of image quality. Images having higher noise
levels do not necessarily undermine diagnostic accuracy; rather, the contrast-tonoise ratio may be similar or improved.
CT dosimetry phantom
Cylinders of polymethylmethacrylate (PMMA) used for standard measurements
of dose in CT, having a diameter of 160 mm (head phantom) or 320 mm (body
phantom) and of appropriate length. The phantoms are constructed with removable inserts parallel to the axis to allow the positioning of a dosimeter at the centre
and 10 mm from the outer surface (periphery).
CT number
Number used to represent the mean x-ray attenuation associated with each elemental area of the CT image. Numbers are normally expressed in terms of Hounsfield units (HU). Measured values of attenuation are transformed into CT
numbers using the Hounsfield scale:
CT number ¼
lmaterial lwater
1000 ðHUÞ
lwater
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where l is the effective linear attenuation coefficient of the measured material relative to water for the utilised x-ray beam. The CT number scale is defined so that
water has a value of 0 HU and air a value of 1000 HU.
Detector array
The entire assembly of detectors, including their interspace material, arranged
along an arc or circumference (depending on scanner technology) of a circle centred on the axis of rotation. Multiple arcs or rows comprise the full array.
Detector width
The distance between the two opposite faces of any single detector in the detector
array.
Diagnostic reference levels
These are used for procedures applying ionising radiation in medical diagnosis to
indicate whether, in routine conditions, the radiation dose from a specified procedure is unusually high or low for that procedure.
Dose descriptor
A measurable parameter, such as CTDIw, CTDIvol or DLP, from which the
effective dose or the organ doses from a CT examination can be estimated, or
by which the radiation dose characteristics of different CT scanners can be
compared.
Dose–length product (DLP)
Please refer to details in Appendix A of this publication.
Effective mAs
The tube current–exposure time product (mAs) divided by the pitch; an estimate
of the mAs per slice.
Effective dose
Please refer to details in Appendix A of this publication.
Gantry
The scanner structure containing the x-ray tube, collimators, and the detector
array.
Geometric efficiency
The percentage of the x-ray beam that is intercepted by the detectors.
Helical CT
A particular technique of scanning in which there is continuous rotation of the
x-ray tube coupled with continuous linear translation of the patient through the
gantry aperture in order to achieve volumetric data acquisition. Also known as
spiral or volume CT.
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Hounsfield unit (HU)
See CT number.
Image thickness
Effective thickness of the tomographic section, as measured by the full width at
half maximum of the sensitivity profile in the centre of the scan field.
Interpolation
A mathematical method of averaging or smoothing images that are displayed as a
larger number of pixels than that from which they were originally constructed.
Longitudinal (z) mA modulation
Longitudinal (z) mA modulation addresses the varying attenuation of the patient
among anatomic regions by varying the mA along the z axis (length) of the patient
(e.g., shoulders versus the abdomen versus the pelvis). The task of z modulation is
to produce relatively uniform noise levels across the various regions of anatomy.
MDCT (multi-detector computed tomography)
MDCT systems are CT scanners with a detector array consisting of more than
a single row of detectors. The ‘multi-detector-row’ configuration of MDCT
scanners refers to the use of multiple detector arrays (rows) in the longitudinal
direction (that is, along the length of the patient). MDCT scanners utilise thirdgeneration CT geometry in which the arc of detectors and the x-ray tube rotate
together. All MDCT scanners use a slip-ring gantry, allowing helical acquisition.
Modulation of x-ray tube current (mA)
A method of managing the dose in CT by automatically adjusting the current to
achieve the selected image quality. May be angular or longitudinal mA modulation.
Multiphase examination
An examination in which two or more sets of images of a defined anatomical area
are taken (e.g., before and after the injection of a contrast medium).
Noise
A fundamental characteristic that is present, to some extent, in all images. Noise
tends to reduce the visibility of structures and objects, especially those that have
relatively low contrast. In medical imaging, the objective is not to eliminate the
noise, but to reduce it to a clinically acceptable level. Noise is the point-to-point
variation in image brightness that does not contain useful information. The magnitude of noise is indicated by the standard deviation of the CT numbers within a
region of interest in the image.
Organ dose
The mean absorbed dose in an organ or tissue, i.e., the total energy imparted to
the organ or tissue divided by the total mass of the organ or tissue.
Overbeaming
The situation when the x-ray beam incident to the patient extends beyond the
active detector area and hence is not used for imaging purposes.
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Overranging
The increase in dose–length product due to the additional rotations required for
the helical (spiral) interpolation algorithm.
Partial volume effect or averaging
The inaccuracy in a CT number caused by the presence of a structure that is only
within part of a slice. Such effects become less important as the slice thickness is
reduced.
Phantom
A device that absorbs or scatters radiation in an equivalent manner to a patient,
utilised to estimate radiation doses and test imaging systems without actually
exposing a patient. A phantom may be an anthropomorphic or a physical test
object.
Pitch / Beam pitch
The ratio of the distance the table travels per x-ray tube rotation to the x-ray
beam width. Pitch conveys the degree of overlap of the radiation beam: a pitch
of 1 indicates contiguous radiation beams, a pitch of less than 1 indicates overlap
of the radiation beams, and a pitch greater than 1 indicates a gap between the
radiation beams.
In SDCT, there is normally no difference between beam width and detector width.
But in MDCT, there is difference between the two. Therefore two terms were proposed: ‘beam pitch’ and ‘detector pitch’. The use of beam pitch (or pitch) as given
above is applicable equally to both SDCT and MDCT.
Polymethylmethacrylate (PMMA)
A commercially available plastic polymer, e.g., Perspex or Lucite.
Radiation effect, deterministic
A radiation effect for which there exists a threshold level of dose. Above the
threshold dose, the severity of the effect is greater as the dose increases.
Radiation effect, stochastic
A radiation effect occurring without a threshold dose, whose probability is proportional to the dose and whose severity is independent of the dose.
Raw data
The values of x-ray detector response from all views and rays within a scan.
Reconstruction algorithm
A mathematical procedure used to convert the collected data into an image. Different algorithms are used to emphasise, enhance, or improve certain aspects of
the data.
Scan time
For a single exposure, the time interval between the beginning and the end of the
acquisition of attenuation data. This may be longer than the exposure time owing
to pulsing of the x-ray emission with some CT scanners.
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SDCT (single-detector computed tomography)
Computed tomography technology that utilises a single row of detectors, permitting scanning of only one slice at a time in either a discrete (sequential) or continuous (spiral) acquisition.
Slice
A tomographic section (defined by the position and thickness) of a test phantom
or patient under investigation during a single CT exposure.
Slice thickness
In this document, the term ‘slice thickness’ refers to that used for data acquisition (slice collimation). Larger slice thicknesses (2.5 mm, 5 mm, 10 mm) can be
generated by electronically combining the signal from several detector rows.
Therefore the slice thickness used for the purposes of image review often differs
from the slice thickness used for data acquisition. It may be larger, but is never
smaller.
Spatial resolution
The ability to resolve different objects in the displayed CT image, when the difference in attenuation between the objects and the background is large compared to
noise; normally requires a difference corresponding to at least one hundred HU.
Spiral CT
See helical CT.
Temporal mA modulation
Temporal mA modulation alters the tube current according to a time-based criterion. This is most commonly used in CT examinations of the heart, reducing the
dose for projections of limited interest, such as in early systole where the rapid
cardiac motion compromises image quality.
Tube current–exposure time product (mAs)
The product of x-ray tube current (mA) and the exposure time in seconds (s).
Volume CT
See helical CT.
X-ray tube voltage
Potential difference applied between cathode and anode of an x-ray tube.
Z modulation
See longitudinal modulation.
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1. MDCT TECHNOLOGY
(1) Modern CT scanners employ multiple rows of detector arrays allowing rapid
scanning and wider scan coverage.
(2) All new CT systems have multiple detectors with a single or dual x-ray source,
and a number of new dose reduction tools have become available commercially.
1.1. Background
(3) Computed Tomography (CT) technology and its clinical applications have
shown enormous resilience against alternative diagnostic methods and, at the moment, the use of CT continues to rise. Current technology provides high power xray tubes, enormous computing power, multichannel detectors to give sub-millimetre
slices with wider scan coverage and faster rotation times to complete one rotation
in one-third of a second. CT now has an important role in dynamic applications such
as cardiology and three-dimensional imaging of vascular and musculoskeletal
anatomy.
(4) A number of terms are used for this technology, namely multi-detector-row
computed tomography (MDRCT), multi-detector CT (MDCT), multi-detector-array
spiral CT, multi-channel CT, and multi-slice CT (MSCT). The number of simultaneous but independent measurements along the patient long axis is often referred
to as the number of ‘slices’, and this value is commonly used to represent the technical
capabilities of a system (e.g., 64-slice MDCT). In this report, the Commission has
chosen to use the terminology MDCT when referring to the technology generically,
and 64-MDCT when referring to a specific technical implementation of MDCT.
(5) In 2000, ICRP published ‘Managing Patient Dose in Computed Tomography’
(ICRP, 2000a). At that time there was an urgent need to focus the attention of radiologists, physicians, medical physicists, and other personnel involved in CT on the
relatively higher doses to organs in individual patients, increasing frequency of CT
examinations, changes in clinical applications, and the increasing contribution of
CT to the collective dose. Further, the technology in use dominantly utilised a single
row of detectors (SDCT), permitting scanning of only a single slice at a time in either
a discrete (sequential acquisition) or continuous fashion (spiral acquisition). Multiple detector rows along the z axis (longitudinal axis of the patient, i.e., head to
toe) permit simultaneous scanning of more than one slice. MDCT was in its infancy
at the time of the 2000 report (ICRP, 2000a) and thus there was only brief mention in
the report of its impact on radiation dose. The data and experience were insufficient
to make any evaluation. In the following years there was a phenomenal increase in
use of MDCT, and technology has been advancing very rapidly to move from 4 slices
to 8, 16, 32, 40, and 64 slices. Furthermore, dual-source CT has been recently made
available and a 256-slice MDCT system will soon be available. The improved speed
of MDCT scanning has also meant new applications (cardiac CT, whole body scanning) as well as improved patient throughput and workflow. In the past two decades,
the use of CT scanning has increased by more than 800% globally (Frush, 2003). In
the United States, over the period of 1991 to 2002, a 10–20% growth per year in CT
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procedures has been documented (Fox, 2003). Also in the United States during this
period, CT scanning for vascular indications has shown a 235% growth, followed by
a 145% growth in cardiac applications. An increase has also been demonstrated in
abdominal (25%), pelvic (27%), thoracic (26%), and head and neck (7%) applications
(Fox, 2003). With 64-slice MDCT, a further substantial increase is expected in cardiac applications. A 10% annual growth in the global CT market was reported in the
year 2002 and this trend seems set to continue.
1.2. Introduction to MDCT Technology
(6) MDCT systems are CT scanners having more than one detector array, or row,
in the arc around the patient. In CT, the detector array consists of 600 to 900 detector elements and corresponds to one transaxial plane. In SDCT, there is one such
row of detectors. The ‘multi-detector-row’ nature of MDCT scanners refers to the
use of multiple detector arrays (rows) in the longitudinal direction (that is, along
the length of the patient lying on the patient table). MDCT scanners in most cases
utilise CT geometry in which the arc of detectors and the x-ray tube rotate together.
All MDCT scanners use a slip-ring gantry, allowing spiral acquisition at rotation
speeds as fast as 0.33 second for a full rotation of 360 degrees of the x-ray tube
around the patient. A scanner with two rows of detectors (Elscint CT Twin) had already been available since 1992 when several manufacturers introduced MDCT
scanners with four detector rows in 1998. The primary advantage of these scanners
is the ability to scan more than one slice simultaneously and hence use the radiation
delivered from the x-ray tube more efficiently (Fig. 1.1). The time required to scan a
given volume could thus be reduced considerably. The number of slices, or data
channels, acquired per axial rotation continues to increase, with 64-detector systems
Fig. 1.1. Schematics of detector elements in the arc around the patient, and detector rows perpendicular to
the z axis. All CT scanners have many detector elements in the arc around the patient. However, only
MDCT systems have more than one row perpendicular to the z axis.
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now being widely available (Flohr et al., 2005a, 2005b). It is likely, in the coming
years, that even larger arrays of detectors having longitudinal coverage per rotation
> 4 cm will be commercially available. Preliminary results from a 256-detector scanner (12.8 cm longitudinal coverage at the centre of rotation) have already been published (Mori et al., 2004). Further, an MDCT system with two x-ray sources is now
commercially available (Flohr et al., 2006), signalling continued evolution of CT
technology and applications.
(7) MDCT scanners can also be used to cover a specific anatomic volume with
thinner slices. This improves the spatial resolution in the longitudinal direction considerably without the drawback of extended scan times. Improved resolution in the
longitudinal direction is of great value in multiplanar reformatting (perpendicular or
oblique to the transaxial plane) and in three-dimensional (3D) representations. Spiral scanning additionally allows overlapping data sets – which improve multiplanar
reconstruction and 3D image quality – to be reconstructed without additional radiation dose to the patient.
1.3. Differences between SDCT and MDCT
(8) One essential difference between SDCT and MDCT is how the thickness represented by an image, or slice, is determined. For SDCT, slice thickness is determined by a combination of pre-patient and post-patient collimation. Therefore,
the dimension of the detector array along the longitudinal axis can extend beyond
the anticipated width of the x-ray beam or image slice (Fig. 1.1), i.e., the detector
width is greater than the beam width. For MDCT, the converse is true and the
x-ray beam width must be large enough to allow irradiation of all ‘active’ detector
rows, i.e., all those being used for a particular scan acquisition; slice thickness is instead determined by the widths of the individual active detector rows used for image
reconstruction.
(9) In Fig. 1.1 the single-detector row CT (SDCT) system on the left has one detector row, which is perpendicular to the longitudinal axis (indicated by z). Within this
one row are many detector elements in the arc around the patient. On the right the
MDCT system has 16 detector rows, which are all perpendicular to the z axis. For
SDCT, the width of the detector (relative to the centre of the gantry) is 20 mm,
although the pre-patient collimated beam is generally no more than 10 mm. Thus,
in SDCT, the detector is wider than the x-ray beam, ensuring that the entire primary
x-ray beam is detected. The multiple-detector-row CT (MDCT) system on the right
has 16 detector elements each 1.25 mm along the longitudinal axis for each of the
approximately 900 positions around the patient. The width of the detector is also
20 mm at the isocentre. Depending on the scanner model and collimation used, this
may result in some portion of the primary radiation beam not being detected. This
overbeaming phenomenon is described in Section 3.2.1.
(10) The four data channels in 4-MDCT allow the acquisition of four simultaneous slices, of 1.25, 2.5, 3.75 or 5 mm width. Wider image thicknesses (2.5 mm, 5
mm, 10 mm) are generated by electronically combining the signal from more than
one of these rows. Therefore the image thickness used for the purposes of image
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review often differs from the slice thickness used for data acquisition. In this
document, the term ‘slice thickness’ always refers to that used for data acquisition
(slice collimation).
(11) Owing to the narrow width of the detector rows and the use of third-generation geometry, gas ionisation detectors are no longer used for MDCT scanners. In
order to generate an image of a 1-mm slice of anatomy, detector rows of not much
more than 1 mm in width must be used (detector dimensions are normalised relative
to their coverage at the centre of the CT gantry).
(12) Another design for an MDCT detector array is illustrated in Fig. 1.2. When
small slices are desired, only the central portion of the array is used. It is therefore
not necessary to have narrow rows in the outer portions of the array. The wider
detectors at the periphery allow simultaneous acquisition of four slices each of 5
mm thickness. This design is somewhat less expensive and geometrically more
efficient.
(13) Currently, MDCT systems are capable of acquiring up to 64 slices simultaneously along the z direction (Fig. 1.3). Three of the four manufacturers use 64 rows
of either 0.625 mm or 0.5 mm detectors. The fourth manufacturer uses 32 rows of 0.6
mm detectors and oscillates the focal spot to acquire 64 overlapping slices (Flohr
et al. 2005b). This results in the reduction of spiral artefacts and improved spatial
resolution along the longitudinal axis (Flohr et al. 2005b).
(14) For sequential or axial data acquisitions (e.g., the table is stationary during
the rotation of the x-ray tube around the patient), each channel collects sufficient
data to create one ‘slice’ or image, so as many as 64 independent images along the
z axis could theoretically be reconstructed. For narrow slice widths, geometrical
‘cone-beam’ considerations may limit the number of allowed images per rotation
to less than 64. For example, one manufacturer’s 16-MDCT scanner allows only
12 data channels to be used in sequential scanning because of cone beam considerations (Flohr et al. 2005a, 2005b).
(15) The primary attribute of MDCT systems is not the number of physical detector rows, but the number of slices that are acquired simultaneously. The speed
needed to cover a given volume is improved by a factor equivalent to the number
of slices included in the scan simultaneously. The reason why the number of simultaneous slices was initially limited to four was the amount of data to be acquired and
Fig. 1.2. Diagram of the detector geometry used in a MDCT from two major manufacturers. The detector
array is 20 mm wide along the longitudinal (z) axis and uses eight detector rows of varying widths to allow
simultaneous scanning of 4 slices up to 5 mm thick.
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Fig. 1.3. Diagram of the detector geometries used in 64-MDCT from four major manufacturers. The
Siemens 64-MDCT uses 32 sub-mm detectors and a moving focal spot to achieve 64 overlapping slice
measurements.
transferred simultaneously. At that time, engineering and cost considerations limited
the systems to four simultaneous data collection systems. Additionally, cone beam
artefacts were not severe in 4-MDCT systems, but as the number of simultaneous
slices increased, these artefacts become more problematic using conventional fanbeam reconstruction methods. Once 3-D cone-beam reconstruction algorithms (or
advanced fan-beam algorithms with cone-beam corrections), and the increased computational power needed for these algorithms, became available, 8- and 16-MDCT
scanners were introduced.
(16) The advent of spiral CT initially introduced an additional acquisition parameter into the CT vocabulary, pitch. Pitch is the ratio of the distance of table travel per
x-ray tube rotation to the x-ray beam width. With MDCT, a significant amount of
confusion was introduced regarding the definition of pitch, as some manufacturers
used an altered definition of pitch that related the distance of table travel per
x-ray tube rotation to the width of an individual data channel, resulting in calculated
pitch values of 3 or 6. The International Electrotechnical Commission CT Safety
Standard re-established the original definition of pitch (distance of table travel normalised to the total beam width) as the only acceptable definition of pitch (IEC 2002,
McCollough and Zink 1999). This definition of pitch conveys the degree of overlap
of the radiation beam: a pitch of 1 indicates contiguous radiation beams, a pitch less
than 1 indicates overlap of the radiation beams, and a pitch greater than 1 indicates
gaps between the radiation beams.
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(17) Two manufacturers report the tube current–exposure time product (milliampere second – mAs) as the average mAs per unit length along the longitudinal axis,
called either effective mAs or mAs/slice, and calculated as actual mAs/pitch. This
distinction between mAs and mAs per unit length is important because, as the pitch
is increased, scanner software may automatically increase the mA such that the image noise (and patient dose) remains constant with increasing pitch values (Flohr
et al. 2003a, 2003b, Mahesh et al. 2001). When the effective mAs or mAs/slice is displayed, the user may be unaware that the actual mA changes when the pitch value is
changed. On other MDCT systems, the mA value is automatically adjusted to the
value that will keep image noise constant as pitch or slice width is changed, and
the selection box is turned orange to alert the user to the change in the prescribed
mA value.
1.4. Upcoming developments
(18) Recently, cone-beam computed tomography (CBCT) has been introduced
that employs a large-area detector with approximately 1000 or more detector
‘rows’ (compared to 16–64 rows in current clinical MDCT systems). CBCT with
one or two flat-panel detectors is a developing imaging modality. There is no true
cone beam commercial system available yet. The available reports from experimental systems indicate its usefulness in intraoperative imaging, interventional
radiology, bone and lung imaging, mammography and radiation therapy (Siewerdsen et al. 2005, Daly et al. 2006, Ross et al. 2006, Guerrero et al. 2006, Glick et al.
2007). The advantages include decreased scan time, wide z-axis coverage and higher near-isotropic spatial resolution. Current weaknesses of the experimental platform arise from lack of integration with an in-room navigation system,
difficulties with scatter rejection, the presence of artefacts, the use of fixed collimation at the output of the x-ray source, and inaccuracy of the dosimetry method
using a standard 100-mm-long CT chamber (Ross et al. 2006, Siewerdsen et al.
2005).
(19) Volumetric imaging of the whole breast has been demonstrated by a number
of investigators using a flat panel detector for CBCT breast imaging (Glick et al.
2007, Kwan et al. 2007, Shaw et al. 2005). The framework for determining the optimal design and combination of parameters was the patient dose criteria, i.e., mean
glandular dose (MGD) constrained to that given for a typical two-view conventional
mammography study (Glick et al. 2007). The application of CBCT in dental and
maxillofacial imaging is also increasing (Guerrero et al. 2006, Sukovic 2003). Some
studies promoting dose reduction have been published (Tsiklakis et al. 2005, Ludlow
et al. 2003).
1.5. What is the motivation for this report?
(20) With ICRP Publication 87 (ICRP, 2000a), an editorial in the British Medical
Journal (Rehani and Berry, 2000), and the February 2001 issue of AJR (AJR, 2001),
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ICRP Publication 102
considerable attention was focused on the topic of dose management in CT. Two
papers addressed the lack of appropriate parameter selection in paediatric CT
examinations (Paterson et al. 2001, Donnelly et al. 2001). Further, Brenner
et al. (2001) reported on the potential risk of cancer induction from the use of
CT in the paediatric population. These publications note that the use of CT has
significantly increased in children (for clinically valid reasons), but they warned
that this increased usage carries with it a potential for excessive exposure to radiation and an increased risk of cancer in the paediatric population. In the editorial
by Lee F. Rogers in the same issue of AJR (Rogers, 2001), he stated: ‘sorry to say,
but kids get overlooked’. These reports aroused media attention and the clinical
and radiological protection communities recognised that radiation doses in CT
should be more carefully scrutinised. The number of publications on radiation
exposure in CT, and management thereof, has been steadily increasing. Manufacturers now put more emphasis on radiation exposure reduction and improved image optimisation in addition to scan time reduction. In recent years, improved
management of radiation dose in CT has been high on the agenda for all CT
manufacturers.
(21) In 2005, the Commission realised that essentially all new CT systems are
MDCT, and that a number of new dose reduction tools have become available commercially. Thus, to address these new tools, the continued growth in CT applications, and the consequent growth in the contribution of CT to the collective dose
from medical applications, it was decided to update ICRP Publication 87 (ICRP,
2000a). In addition to reviewing these technology changes in CT, a number of issues
will be addressed, such as:
Has MDCT caused an increase or decrease in patient doses?
In cases where patient doses have increased, why is this so?
How does new technology contribute to dose management?
What actions are required of the scanner operator?
Are there specific educational actions still required?
(22) As in its previous report (ICRP 2000a), the primary audience for this document is imaging professionals: radiologists, cardiologists, operators, medical physicists and researchers involved in patient dose management. However, this
document provides reference material that may be useful for referring physicians,
physicians who may own a CT scanner, regulators and national authorities, manufacturers and hospital administrators.
1.6. References for Chapter 1
AJR, 2001. Am. J. Roentgenol. 176, 287–306.
Brenner, D., Elliston, C., Hall, E., et al., 2001. Estimated risks of radiation-induced fatal cancer from
paediatric CT. Am. J. Roentgenol. 176, 289–296.
Daly, M.J., Siewerdsen, J.H., Moseley, D.J., et al., 2006. Intraoperative cone-beam CT for guidance of
head and neck surgery: Assessment of dose and image quality using C-arm prototype. Med. Phys. 33
(10), 3767–3780.
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Donnelly, L.F., Emery, K.H., Brody, A.S., et al., 2001. Minimizing radiation dose for paediatric body
applications of single-detector helical CT: Strategies at a large children’s hospital. Am. J. Roentgenol.
176, 303–306.
Flohr, T., Ohnesorge, B., Bruder, H., et al., 2003a. Image reconstruction and performance
evaluation for ECG-gated spiral scanning with a 16-slice CT system. Med. Phys. 30, 2650–
2662.
Flohr, T., Stierstorfer, K., Bruder, H., et al., 2003b. Image reconstruction and image quality evaluation
for a 16-slice CT scanner. Med. Phys. 30, 832–845.
Flohr, T.G., Schaller, S., Stierstorfer, K., et al., 2005a. Multi-detector row CT systems and imagereconstruction techniques. Radiology 235, 756–773.
Flohr, T.G., Stierstorfer, K., Ulzheimer, S., et al., 2005b. Image reconstruction and image
quality evaluation for a 64-slice CT scanner with z-flying focal spot. Med. Phys. 32, 2536–
2547.
Flohr, T.G., McCollough, C.H., Bruder, H., et al., 2006. First performance evaluation of a dual-source
CT (DSCT) system. Eur. Radiol. 16, 256–268.
Fox, S.H., 2003. Emerging developments in multi-detector CT. Presented at Advances in Multidetector
CT Meeting, Washington, DC, September 13–14.
Frush, D.P., 2003. CT radiation dose optimisation in children. Presented at Advances in Multi-detector
CT Meeting, Washington, DC, September 13–14.
Glick, S.J., Thacker, S., Gong, X., et al., 2007. Evaluating the impact of x-ray spectral shape on image
quality in flat-panel CT breast imaging. Med. Phys. 34 (1), 5–24.
Guerrero, M.E., Jacobs, R., Loubele, M., et al., 2006. State-of-the-art on cone beam CT imaging for preoperative planning of implant placement. Clinical Oral Investigations 10 (1), 1–7.
ICRP, 2000a. Managing patient dose in computed tomography. ICRP Publication 87, Ann. ICRP
30(4).
IEC, 2002. Medical Electrical Equipment. Part 2-44: Particular requirements for the safety of x-ray
equipment for computed tomography. IEC publication No. 60601-2-44. Ed. 2.1. International
Electrotechnical Commission (IEC) Central Office, Geneva, Switzerland.
Kwan, A.L., Boone, J.M., Yang, K., et al., 2007. Evaluation of the spatial resolution characteristics of a
cone-beam breast CT scanner. Med. Phys. 34 (1), 275–281.
Ludlow, J.B., Davies-Ludlow, L.E., Brooks, S.J., 2003. Dosimetry of two extraoral direct digital imaging
devices: NewTom cone beam CT and Orthophos Plus DS panoramic. Dentomaxillofacial Radiol. 32
(4), 229–234.
Mahesh, M., Scatarige, J.C., Cooper, J., et al., 2001. Dose and pitch relationship for a particular multislice CT scanner. Am. J. Roentgenol. 177, 1273–1275.
McCollough, C.H., Zink, F.E., 1999. Performance evaluation of a multi-slice CT system. Med. Phys. 26,
2223–2230.
Mori, S., Endo, M., Tsunoo, T., et al., 2004. Physical performance evaluation of a 256-slice CT-scanner
for four-dimensional imaging. Med. Phys. 31, 1348–1356.
Paterson, A., Frush, D.P., Donnelly, L.F., 2001. Helical CT of the body: Are settings adjusted for
paediatric patients? Am. J. Roentgenol. 176, 297–301.
Rehani, M.M., Berry, M., 2000. Radiation doses in computed tomography. The increasing doses of
radiation need to be controlled (Editorial). BMJ 4; 320, 593–594.
Rogers, L.F., 2001. Taking care of children. Check out the parameters used for helical CT (Editorial). Am.
J. Roentgenol. 176, 287.
Ross, W., Cody, D.D., Hazle, J.D., 2006. Design and performance characteristics of a digital flat-panel
computed tomography system. Med. Phys. 33 (6), 1888–1901.
Shaw, C., Chen, L., Altunbas, M., et al., 2005. Cone beam breast CT with a flat panel detector –
Simulation, implementation and demonstration. Conf. Proc. IEEE Eng. Med. Biol. Soc. 4, 4461–
4464.
Siewerdsen, M.J., Moseley, D.J., Burch, S., et al., 2005. Volume CT with a flat-panel detector on a mobile,
isocentric C-arm: Pre-clinical investigation in guidance of minimally invasive surgery. Med. Phys. 32
(1), 241–254.
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Sukovic, P., 2003. Cone beam computed tomography in craniofacial imaging. Orthod. Craniofac. Res. 6
(Suppl. 1), 31–36.
Tsiklakis, K., Donat, C., Gavala, S., et al., 2005. Dose reduction in maxillofacial imaging using low dose
cone beam CT. Eur. J. Radiol. 56 (3), 413–417.
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