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Am J Physiol Heart Circ Physiol 301: H269–H278, 2011.
First published May 13, 2011; doi:10.1152/ajpheart.00320.2011.
Review
Doppler velocity measurements from large and small arteries of mice
Craig J. Hartley,1 Anilkumar K. Reddy,1 Sridhar Madala,2 Mark L. Entman,1 Lloyd H. Michael,1
and George E. Taffet1
1
Section of Cardiovascular Research, Department of Medicine, Baylor College of Medicine, and the Methodist DeBakey
Heart and Vascular Center, Houston; and 2Indus Instruments, Webster, Texas
Submitted 30 March 2011; accepted in final form 7 May 2011
peripheral vascular flow; pulse wave velocity; coronary flow reserve; transverse
aortic constriction; pressure overload hypertrophy
THIS REVIEW ARTICLE is part of a collection on Assessing Cardiovascular Function in Mice: New Developments and
Methods. Other articles appearing in this collection, as well as
a full archive of all Review collections, can be found online at
http://ajpheart.physiology.org/.
The ability to alter the genotype of the mouse has produced
numerous models for studying cardiovascular physiology and
pathophysiology and has generated a need to evaluate the
changes that occur in mice during phenotypic development and
maturation (8). The genetic manipulations can alter the structure, anatomy, pathology, and physiology of cells, organs, or
systems in ways that can be subtle and unpredictable and that
often change with time (6, 33). The major problems in adapting
existing methods for use in mice relate to the small size and
high heart rates that place extreme demands on both spatial and
temporal resolution (10). We will focus here on noninvasive
ultrasonic Doppler methods, which our group and others have
developed and used to evaluate cardiovascular physiology and
Address for reprint requests and other correspondence: C. J. Hartley, Dept.
of Medicine (CVS), Baylor College of Medicine, One Baylor Plaza, Houston,
TX, 77030 (e-mail: [email protected]).
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function in anesthetized mice and which can be used serially to
follow changes because of aging, remodeling, and/or the effects of surgical or pharmacological interventions (19, 66). The
utility and value of the measurements will be illustrated using
the widely used model of transverse aortic constriction (TAC)
that was developed for producing pressure-overload cardiac
hypertrophy and remodeling (51) but which produces numerous other measurable changes in peripheral vascular mechanics
and function (23, 34).
Murine Doppler System
One of the most promising noninvasive methods available
for use in mice is ultrasound, and Doppler methods are particularly attractive for sensing blood flow velocity in the heart and
major blood vessels (19). Doppler instruments measure blood
velocity by detecting the difference in frequency between an
emitted burst of ultrasound and the returning echoes from
moving blood. In a pulsed-Doppler system the sample volume
can be adjusted in depth by varying the time delay between
transmission of a short ultrasonic burst and sampling of the
returning echoes. The resulting Doppler signal, which is usually in the audible range, is a summation of the Doppler-shifted
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Hartley CJ, Reddy AK, Madala S, Entman ML, Michael LH, Taffet GE.
Doppler velocity measurements from large and small arteries of mice. Am J Physiol
Heart Circ Physiol 301: H269 –H278, 2011. First published May 13, 2011;
doi:10.1152/ajpheart.00320.2011.—With the growth of genetic engineering, mice
have become increasingly common as models of human diseases, and this has
stimulated the development of techniques to assess the murine cardiovascular
system. Our group has developed nonimaging and dedicated Doppler techniques for
measuring blood velocity in the large and small peripheral arteries of anesthetized
mice. We translated technology originally designed for human vessels for use in
smaller mouse vessels at higher heart rates by using higher ultrasonic frequencies,
smaller transducers, and higher-speed signal processing. With these methods one
can measure cardiac filling and ejection velocities, velocity pulse arrival times for
determining pulse wave velocity, peripheral blood velocity and vessel wall motion
waveforms, jet velocities for the calculation of the pressure drop across stenoses,
and left main coronary velocity for the estimation of coronary flow reserve. These
noninvasive methods are convenient and easy to apply, but care must be taken in
interpreting measurements due to Doppler sample volume size and angle of
incidence. Doppler methods have been used to characterize and evaluate numerous
cardiovascular phenotypes in mice and have been particularly useful in evaluating
the cardiac and vascular remodeling that occur following transverse aortic constriction. Although duplex ultrasonic echo-Doppler instruments are being applied to
mice, dedicated Doppler systems are more suitable for some applications. The
magnitudes and waveforms of blood velocities from both cardiac and peripheral
sites are similar in mice and humans, such that much of what is learned using
Doppler technology in mice may be translated back to humans.
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echoes from many blood cells moving at different velocities,
and the spectrum of frequencies represents the distribution of
red cell velocities within the sample volume. The following
equation relates the Doppler shift frequency (⌬f) to the velocity
(V) of a reflector:
⌬f ⫽ 2f o(V ⁄ c)cos␪
(1)
Ultrasonic Doppler methods can also be used to measure
cardiac motion (17, 31, 55) and artery diameter pulsation (20)
in mice, which when combined with pressure provide an index
of vessel compliance (29). Doppler devices actually measure
the phase of blood echoes that advance (or recede) continuously to generate the Doppler shift frequency (3). Echoes from
slower moving solid structures such as blood vessel walls that
move back and forth generate phase signals which advance and
recede with each cardiac cycle. One cycle of echo phase (360°)
corresponds to reflector motion of 1/2 wavelength (because of
the 2-way path) or about 38 ␮m at 20 MHz (18). Under optimal
conditions, Doppler methods can resolve about 1° of phase or
0.1 ␮m of tissue motion and can display cardiac or vessel wall
motion continuously throughout the cardiac cycle (20).
Examples and Applications
Although spectral (57), tissue (31, 45), color (57), and power
(32) Doppler modalities are available and have been used to
assess cardiovascular function and flow in mice, we will
concentrate here on spectral Doppler measurements of blood
flow velocity and vessel wall motion in peripheral vessels of
anesthetized mice. Blood flow velocity can be measured from
the left ventricular inflow and outflow tracts and from peripheral arteries smaller than ⬃100 ␮m diameter. The protocols
from which the following examples were obtained were approved by the Institutional Animal Care and Use Committee of
Baylor College of Medicine and conform with the Guide for
the Care and Use of Laboratory Animals, published by the
National Institutes of Health (NIH Publication No. 85-23,
Revised 1996).
An example of cardiac Doppler signals from a mouse using
a 10-MHz transducer placed at the base of the sternum and
pointed toward the heart is shown to the right in Fig. 2 along
with the ECG. A photo of a mouse on an ECG/heater board is
shown to the left. From this approach cardiac timing and
several indexes of left ventricular systolic function (peak velocity and acceleration), diastolic function (peak-early and
late-filling velocities and their ratio), and the Tei index (55) can
be measured (59). In this example, which was done to illustrate
Fig. 1. Drawing and photos of a 20-MHz Doppler
probe designed to measure blood flow velocity noninvasively in mice. The probe consists of a ⬃1-mm
square piezoelectric crystal, air-backed using Styrofoam, and mounted into a 2-mm-diameter stainlesssteel tube with epoxy. A concave epoxy lens is molded
to the front face to focus the sound beam at a depth of
⬃4 mm. A 10-MHz probe uses a 3-mm tube and is
focused at ⬃6 mm.
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where fo is the ultrasonic frequency (10 or 20 MHz for use in
mice), c is the speed of sound in blood (⬃1,540 m/s), and ␪ is
the angle between the sound beam and the true direction of
flow or motion. When the sound beam is parallel to the
direction of flow (␪ is zero), the conversion factors relating
velocity to the Doppler shift are 7.5 (cm/s)/kHz at 10 MHz and
3.75 (cm/s)/kHz at 20 MHz (3). The murine pulsed-Doppler
system was adapted from nonimaging technology originally
designed by our group for use with catheter-tip and implantable
probes to measure blood flow in small vessels of humans and
larger animals (16). For noninvasive applications in mice,
small handheld probes and a high-fidelity spectrum analyzer
and signal processor were designed. The probes consist of a
⬃1 mm square of 10- or 20-MHz piezoelectric material
mounted at the end of a 2-mm-diameter, 10-cm-long stainless
steel tube as shown in Fig. 1. A concave epoxy lens is molded
to the front face of the piezoelectric element to focus the sound
beam at a depth of 3– 6 mm. The resulting sample volume is
⬍0.02 ␮l (0.3 mm diameter ⫻ 0.3 mm long) at the focal point.
The Doppler signal processor is a computer-based system that
was designed in collaboration with Indus Instruments, Webster, TX (46, 49). The current generation system digitizes and
stores the audio Doppler signals at 125 kHz from up to four
probes, generates and displays a complex fast Fourier transform in real time, and also captures and displays auxiliary
signals such as ECG and pressure. Imaging is not included. The
Indus software can also calculate a waveform of the spectral
peak velocity for export along with the auxiliary signals for
further processing by other analysis packages such as Excel.
The best velocity resolution is 5 mm/s at 20 MHz, the maximum measurable velocity is 9.3 m/s at 10 MHz, and the best
temporal resolution is 0.1 ms (46). Most ultrasonic imaging
systems also have spectral Doppler capability, and many of the
measurements illustrated here can be made with those systems.
Doppler Displacement
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the technique, the relatively large sample volume included both
the inflow and outflow tracts of the left ventricle. With a
focused probe and proper positioning, signals from the inflow
tract and the aortic root can be isolated. Although Doppler
signals can provide measures of cardiac timing and function
(41, 55), imaging of cardiac structures and their motion is often
required to put the velocity indexes into proper context in the
estimation of ventricular function (57).
With the use of a 20-MHz probe, Doppler signals can be
obtained from many peripheral arterial sites as shown in Fig. 3.
All of the signals except for coronary flow were obtained from
the same mouse. By the measurement of peak and mean
velocity and the evaluation of the shapes of the waveforms
from several arterial sites, changes in regional peripheral vascular impedance can be detected (49). Other groups have used
duplex echo-Doppler systems to measure spectral Doppler
velocity signals from the aortic arch (12), the renal artery (4,
58), the hepatic artery (5), the coronary artery (54), and the
umbilical artery (38) of anesthetized mice.
Pulse Wave Velocity
One of the applications for peripheral measurements of
blood velocity is the determination of arterial pulse wave
velocity (PWV) as an index of arterial stiffness (2, 68). Stiffer
vessels propagate pressure and velocity waves faster than more
compliant vessels. The vertical lines in Fig. 3 are aligned with
the R wave of the ECG, and it can be seen that the time to the
upstroke of velocity increases in proportion to the distance
from the heart. By recording velocity signals from two sites,
separated by a known distance, and measuring the difference in
pulse arrival times, one can determine the pulse transit time
and calculate PWV (21, 24). Velocity can be measured sequentially from the two sites with a high-fidelity ECG used as a
Fig. 3. Doppler velocity signals from several peripheral arterial
sites in an anesthetized mouse using a 2-mm-diameter 20-MHz
probe. All signals were obtained with the mouse supine except
for the renal signals, which were obtained with the mouse prone
and the probe placed lateral to the spine. The coronary signals
were obtained later from a second mouse. The vertical lines
correspond to the R wave of the ECG and are used to measure
velocity pulse arrival times for calculation of pulse wave
velocity (PWV) [from Hartley, et al. (25)].
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Fig. 2. Illustration of the setup used to make noninvasive ECG and Doppler measurements showing an anesthetized mouse taped to electrodes on a
temperature-controlled circuit board, Doppler and ECG signal processors, and a display of ECG and Doppler signals from the left ventricular inflow and outflow
tracts. The 23 ⫻ 30-cm-printed circuit board contains 4 stainless-steel ECG electrodes, an array of 50 surface-mount resistors, a temperature sensor, a large
ground plane, electrical connections to an ECG amplifier, and a temperature controller. Vel, velocity; ⌬f, change in frequency. Labeled on the Doppler tracing
are the opening (o) and closing (c) of the mitral (m) and aortic (a) valves, peak ejection velocity (P) and acceleration (Accel), and peak-early (E) and late-filling
(A) velocities. From these signals, one can obtain accurate timing of cardiac events such as preejection time, filling and ejection times, and isovolumic contraction
and relaxation times as indexes of systolic and diastolic ventricular function [from Hartley et al. (25)].
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with a single probe are used, any variability in the preejection
time or changes in PWV between the velocity measurements
will also add to the measurement error. Consistency in measurement sites and probe orientations minimizes the variability
in PWV measurements and maximizes the ability to detect
differences in groups of mice.
Pressure Drop at a Stenosis
A common method for stressing the cardiovascular system
in wild-type and mutant mice is a chronic pressure overload
model produced by TAC between the origins of the innominate
and left carotid arteries (51). This model produces cardiac
hypertrophy and heart failure in 1– 4 wk, but it is difficult to
measure the degree of stenosis or the pressure drop or to
predict the resulting hypertrophy. At the time of death, one can
cannulate (and occlude) both carotid arteries to measure the
pressure difference, but this act is expected to alter flow (and
the pressure drop) significantly. We have found that it is
possible to measure the velocity of the jet produced by the
stenosis and employ the simplified Bernoulli equation developed by cardiologists (26) to estimate the pressure drop:
⌬P ⫽ 4V2
where ⌬P is the pressure drop (in mmHg) and V is the peak jet
velocity (in m/s). We validated this equation in mice as
illustrated in Fig. 4 by cannulating both carotid arteries in five
mice after TAC and measuring the pressures and the jet
velocities as the pressure drops and flows were varied over a
large range (35). The mouse shown in Fig. 4 had an arrhythmia
such that the jet velocities and pressure drops changed with
each cardiac cycle and each cardiac cycle produced one data
point on the graph. There is an excellent agreement (R2 ⫽ 0.94)
between the measured and calculated pressure drops.
Practical considerations and pitfalls. When the pressure
drop across a stenosis estimated, it is important to align the
sound beam parallel to the direction of flow and to search for
the maximum velocity signal that is usually found just distal to
the stenosis. The Doppler system must be able to resolve
velocities as high as 5 m/s, and because the flow distal to the
stenosis is often turbulent with velocity components traveling
in many directions at once, it is not advisable to use a
nonparallel probe orientation with angle correction to estimate
the peak jet velocity.
Fig. 4. Velocity signals from the high-velocity jet in
a mouse with a transverse aortic constriction (TAC)
along with pressure signals from the right (RCPr)
and left (LCPr) carotid arteries. The mouse shown
had an arrhythmia such that the jet velocity (V) and
the pressure drop (⌬P) varied with each cardiac
cycle. The graph to the right shows data from the 5
pulses shown (including the very small one near the
center) plus data from 4 other mice where the
velocity was varied over a large range. SV, sample
volume. As shown, the maximum jet velocity (in
m/s) measured noninvasively can be used with the
simplified Bernoulli equation (4V2) to estimate the
maximum pressure drop (in mmHg) across the aortic constriction [from Li et al. (35)].
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timing reference, or velocity can be measured from two sites
simultaneously using two Doppler probes (48). Typically, we
determine aortic PWV from velocities measured in the aortic
arch and in the descending aorta about 4 cm distal. Aortic arch
velocity signals are obtained by placing a focused 20-MHz
Doppler probe to the right of the upper sternum and angling the
sound beam down and to the left to align with the direction of
the aortic arch and looking for “aortic shape” velocity signals
moving away from the probe. The sample volume depth is
typically set to ⬃3.0 mm. After recording aortic arch signals,
the probe position is marked and another mark is placed 4 cm
distal on the abdomen. At this location the probe is angled
toward the head and moved laterally looking for strong signals
moving toward the probe. After the signals are recorded and
stored, the fast Fourier transform display is adjusted to maximize temporal resolution (0.1 ms) and the time delay from the
R wave of the ECG to the upstroke of velocity is measured and
averaged over several cardiac cycles at each site. Aortic PWV
(on the order of 4 m/s in mice) (24) is then calculated by
dividing the distance (4 cm) by the difference in pulse arrival
times (on the order of 10 ms) at each site. In previous studies,
we have found that apolipoprotein E-deficient (ApoE⫺/⫺) mice
have elevated PWV compared with age-matched wild-type
mice (21, 62), but in other models the changes are not always
apparent during rest. For instance, in ␣-smooth muscle actin
knockout (␣-SMA⫺/⫺) mice (56), the resting PWV is only
slightly lower than in wild-type mice. However, the response to
a bolus intravenous injection of phenylephrine, which doubles
PWV (from 4.6 to 9.9 m/s) in wild-type mice, is absent in
␣-SMA⫺/⫺ mice (49). In contrast, matrix GLA protein knockout mice (37), which have calcified arteries, have significantly
elevated PWV (10.4 m/s) even at rest (25).
Practical considerations and pitfalls. When PWV is calculated, the ability to accurately time the upstroke of velocity
with respect to the R wave of the ECG is more important than
measuring the magnitude of velocity, and this requires clean
waveforms for both velocity and ECG. A major source of error
is in measuring the distance between the measurement sites
along the curved path of the aorta. This requires careful
measurements of the probe locations along the skin surface, an
estimate of the sample volume positions based on the probe
angles and sample volume depths, and an estimate of the
curved path length. When sequential velocity measurements
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Fig. 5. Baseline and hyperemic left main coronary velocity
signals taken noninvasively from a mouse before, 1 day, and 21
days after TAC. The relative contributions of systolic (S) and
diastolic (D) flow vary between baseline and hyperemia and
with time after constriction [from Hartley et al. (23)].
As shown in Fig. 3, it is also possible to measure left main
coronary flow velocity in mice (13, 15, 52, 53, 66). However,
coronary flow velocity by itself is not as useful or diagnostic
as coronary flow reserve (CFR), which is the ratio of hyperemic flow during maximal vasodilation to baseline resting flow
(H/B) (7, 14, 30). In humans and other animals, adenosine
administered intravenously is often used as a maximal coronary vasodilator (39, 67). While we were attempting to develop
methods to administer adenosine effectively to mice, we discovered that isoflurane, which we use as an anesthetic, had an
effect on coronary flow similar to that of adenosine (22, 50) but
with a smaller change in heart rate. Our protocol for measuring
CFR in mice is to place the mouse anesthetized with 2%
isoflurane in oxygen on the procedure board to which a
micromanipulator has been attached to hold the Doppler probe.
The Doppler probe is placed horizontally to the left of the
sternum with the sample volume depth set at 2 to 3 mm and the
sound beam pointing toward the aortic root and the origin of
the left main coronary artery (22). In the absence of imaging,
coronary velocity signals are identified as flow moving toward
the probe during diastole with a characteristic shape. The probe
holder is then clamped into position, and the sample volume
and probe positions are adjusted to maximize the magnitude,
intensity, and stability of the signal. The isoflurane level is then
set to 1% to obtain and record a baseline level of coronary
velocity and then increased to 2.5% to obtain a hyperemic level
of coronary velocity (9). The sample volume position is reoptimized at each level of anesthesia, and it typically takes about
2 min for coronary velocity to stabilize at each level.
Using the above procedure, we estimated left main CFR in
10 mice before and at several time points for 21 days after TAC
(23). Figure 5 shows waveforms at baseline and at hyperemia
from one mouse before, 1 day after, and 21 days after TAC. In
addition to the expected decrease in the amount of hyperemia
at 21 days, there is also a significant increase in the amount of
flow that occurs during systole. We calculated the hyperemic/
baseline (H/B) ratio from the temporal average spectral peak
velocity at baseline and hyperemia and the systolic-to-diastolic
ratio (S/D) from the areas under the systolic and diastolic parts
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of the spectral peak velocity as shown in Fig. 5. Figure 6 shows
a summary of H/B and S/D at baseline for all mice at the five
time points shown. We found that CFR as estimated by H/B
progressively decreases as the heart adapts and hypertrophies
after TAC and that the amount of flow during systole progressively increases after TAC. After 21 days, CFR is nearly
abolished (1.1), and the amount of left main coronary flow
occurring during systole is nearly equal to that occurring
during diastole (S/D ⫽ 0.92).
We have found that left main CFR as estimated by isoflurane
increases from 2.4 at 6 wk to 3.6 at 2 yr of age but is reduced
from 3.6 to 2.5 in ApoE⫺/⫺ (atherosclerotic) mice (22), from
3.2 to 2.0 in mice with chronic left anterior descending coronary artery occlusions, and from 3.0 to 1.4 in mice with chronic
infusions of angiotensin. The systolic component of coronary
flow is more difficult to measure and is often ignored by
coronary physiologists because systolic flow cannot perfuse
actively contracting myocardium (54, 61). Thus the significant
increase in systolic coronary flow following TAC may indicate
a redistribution of flow away from contracting myocardium or
Fig. 6. Hyperemic/baseline (H/B) coronary velocity and systolic/diastolic
(S/D) area ratios at 5 time points before and after TAC. H/B is an index of
coronary flow reserve and is progressively reduced and nearly abolished (1.1)
after 21 days of aortic constriction. Similarly, the amount of coronary flow that
occurs during systole increases progressively and significantly after aortic
constriction.
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Coronary Flow Reserve
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Carotid Velocities Diameter Pulsations
We find that the pulsations in flow velocity are increased in
the right carotid artery proximal to the stenosis and decreased
in the left carotid artery distal to the stenosis after TAC as
shown in Fig. 7. Also shown in Fig. 7 are carotid artery
diameter signals measured before and after TAC using the
same probe held at a 60° angle to the vessel (20). The signals
were obtained using a modified Doppler module with two
additional sample volumes that were designed to measure
tissue displacement rather than blood velocity and that could be
placed over the near and far vessel walls to measure their
motion during the cardiac cycle. The diameter waveforms
shown were calculated by subtracting the displacement signals
from the near and far vessel walls. Before TAC, the peak
velocities and diameter pulsations are similar in the right and
left carotid arteries. One hour after TAC, both the blood
velocity and diameter pulsations are increased in the right
carotid artery and decreased in the left carotid artery. However,
the mean velocities in each vessel before and after TAC are not
significantly different after 1 h. In a normal mouse (or human)
the total arterial compliance must be sufficient to absorb one
stroke volume per cardiac cycle or the heart cannot eject blood
fully. After TAC, this task must be accomplished by the
ascending aorta and the innominate artery and its branches,
which includes the right carotid artery either by increasing
diameter pulsations or by increasing heart rate, which decreases stroke volume. The average diameter of the right
carotid artery also increases as shown in Fig. 7 such that the
fourfold increase in diameter pulsations operating on a larger
vessel results in a significant increase (⬃8⫻) in volume pulsations. The ability of the vessels originating proximal to the
TAC to remodel and increase compliance is one of the reasons
that mice can survive TAC but cannot survive a more proximal
aortic constriction (51).
Practical considerations and pitfalls. When using tissue
Doppler to measure diameter pulsations, only the pulsations
and not the absolute diameter can be sensed with precision
(⬃0.1 ␮m) (20). The absolute diameter can be estimated from
the distance between the sample volumes positioned over the
near and far artery walls, but the accuracy is poor given the
⬃300 ␮m length of the sample volumes. The absolute diameter
and the magnitude, but not a high-fidelity waveform, of arterial
pulsations can be estimated with a high-resolution (VisualSonics) imaging system (68).
Calculation of Vascular Indexes
In many of the murine models we use, including TAC,
PWV, vascular mechanics, and vascular impedance are altered
in measurable ways, and velocity and diameter measurements
allow for the calculation of several indexes of vascular function
(40). When the heart generates pressure and ejects blood, the
impulse created by the heart propagates into the arterial system
as a forward-traveling wave, and whenever this wave passes a
branch or other discontinuity in vascular impedance, part of the
Fig. 7. Velocity (in cm/s) and diameter change (in ␮m)
signals taken noninvasively from the right and left
carotid arteries of a mouse before and 1 h after TAC.
After TAC, the pulsations in both velocity and diameter
are increased in the right carotid artery proximal to the
band and decreased in the left carotid artery distal to the
band. The photo taken at death 2 wk later shows dilation
of the right carotid artery because of increased pressure
and/or wall shear stress. VP, peak velocity.
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from endocardium to epicardium as the heart hypertrophies
(27, 28, 60). CFR is relatively easy to measure in mice, appears
to be reduced by most forms of heart disease, and could be
used in place of ejection fraction as an index of global cardiac
reserve in mice. The coronary S/D ratio, which does not require
the use of a vasodilator, may also be a useful index of
myocardial perfusion status.
Practical considerations and pitfalls. The Doppler sample
volume is slightly larger than the left main coronary artery, but
because of cardiac motion, the spectral intensity often varies
slightly during the cardiac cycle as shown in Fig. 3. In the
absence of imaging, the proper probe position is verified by
temporarily advancing the sample volume to check for the
presence of aortic velocity signals. Proper measurements of
CFR require that the baseline flow be “normal,” that the
coronary vasodilator be maximal with no other cardiovascular
effects, and that the diameter of the proximal coronary artery
where velocity is sensed be constant. It is unlikely that any of
these conditions are actually met using isoflurane or even
adenosine (67), but it is also unlikely for coronary flow to go
above maximal or below minimal values during estimations of
CFR. The result would be an underestimation of CFR. Nevertheless, our estimates of CFR using isoflurane are similar to
those reported in humans (39) and are higher than those
reported by several other groups in mice using adenosine as a
coronary vasodilator and isoflurane as an anesthetic agent
(52, 67).
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wave is reflected back toward the heart (65). If there were no
reflections, pressure and flow waveforms would be identical,
related by the characteristic impedance of the vessel, and travel
at the PWV. The pressure measured at any site is the sum of
forward and backward pressure waves, and the flow measured
at any site is the difference in the forward and backward flow
waves. This gives us enough information to calculate forward
and backward waves from the measured pressure and flow
waves and the characteristic impedance of the vessel (65),
and if we are allowed to use diameter in place of pressure
and velocity in place of flow, this can be done from the
measured Doppler signals (43). Figure 8 shows carotid
velocity and diameter signals and calculated wave intensity
(43) and forward and backward waves (65) from the common
carotid artery of a mouse. Wave intensity is calculated by
multiplying the derivatives of pressure and flow (or diameter
and velocity). Although the units are wrong when using diameter and velocity, the shapes of the waves and their relative
magnitudes are largely correct (43). The first positive peak in
intensity is the forward-traveling acceleration wave generated
at the opening of the aortic valve, and the second positive peak
is the forward-traveling deceleration wave generated by the
General Discussion
We and others typically use the spectral peak or maximum
velocity waveform because it is the most robust parameter in
the Doppler signal (11). Thus the measured peak velocity
represents the highest component of velocity within the sample
volume in the direction of the sound beam. If one knows the
true direction of flow (only reliable in a straight, unbranching
vessel) and the direction of the sound beam, angle correction
can be used to estimate the true velocity.
Where possible, we always align the Doppler transducer
such that the sound beam is parallel to the axis of the vessel
(and the normal direction of flow) and the Doppler angle (␪) is
close zero. At 0°, a 10% error in angle produces a 1.6% error
in velocity, whereas at 60°, a 10% angle error produces a 32%
Fig. 9. Velocity, diameter change, and forward and
backward waves from the right carotid artery of a
mouse before and after TAC. Note that the scales are
different before and after constriction and that the
second peak seen in the forward wave is missing
after constriction.
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Fig. 8. Doppler velocity and diameter change signals from the carotid artery of
a mouse showing some of the sophisticated analyses that can be done. Wave
intensity (WI) is calculated from the product of the derivatives of pressure and
flow (or diameter and velocity) and indicates the net magnitude and direction
of energy transfer along the vessel. Forward and backward-traveling waves are
calculated from the measured pressure and flow (or diameter and velocity)
waves and can be used to estimate the magnitude and timing of wave
reflections.
closing of the aortic valve. The first negative intensity
waves occur because of reflected waves that decelerate flow
while pressure is still increasing. The shapes of the velocity,
diameter, and wave intensity signals from mice as shown in
Fig. 8 look remarkably similar to those measured in humans
using similar methods (43).
The backward wave is the sum of the reflections from all
sites including the terminal capillaries distal to the measurements site (65). The initial part of the forward wave is generated by the heart, but the timing of the second peak during late
systole suggests that it is caused by the reflected wave from the
descending aorta returning to the heart and then traveling
forward in the carotid artery. We hypothesized that TAC,
which adds significant resistance, would reduce or eliminate
the second peak in the forward wave. Figure 9 shows diameter,
velocity, and forward and backward waves in the right carotid
artery of a mouse before and after TAC. Note that the scale on
the right after TAC is 2.5 times the scale on the left before
TAC and that the second peak in the forward wave before TAC
is indeed eliminated after TAC. Although the ratio of the
magnitudes of backward to forward waves are similar before
and after TAC, the delay to the start of the backward wave is
a little shorter after TAC, suggesting that PWV is higher. We
have not attempted to do a full analysis of the changes in
vascular indexes before and after TAC or in mice with other
vascular abnormalities, but these examples show the potential
of Doppler methodology in quantifying peripheral vascular
mechanics in mice and the similarity of the signals to those
recorded by others from comparable but larger vessels in
humans.
Review
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DOPPLER MEASUREMENTS IN MICE
of 120 kHz is ⬃3 m/s. Imaging is important and necessary
when measuring velocity in cardiac chambers or near the heart
where there are many sites with complex velocity patterns, but
in locations where imaging is not required, dedicated Doppler
systems can employ more optimal ultrasonic parameters and
smaller transducers.
In addition to the spectral Doppler methods illustrated here
to measure peripheral and cardiac blood velocities in mice,
other Doppler modalities are also being applied to mice (57).
Tissue Doppler and strain imaging have been used to measure
cardiac timing (55) and ventricular function (31), color Doppler is used to locate the direction and timing of flow across
valves and in cardiac chambers (57), and power Doppler with
contrast agents is used to assess perfusion in organs such as the
brain and kidney (58) and in tumors (32).
Duplex Doppler/Imaging Systems
DISCLOSURES
Many of the peripheral vascular Doppler measurements
illustrated here can be made with clinical or duplex echoDoppler systems (4, 5, 52) or with one of the VisualSonics
mouse ultrasonic systems (1, 12, 58). However, the optimal
ultrasonic frequencies, probe configurations, and display modalities for imaging and Doppler are different such that some
trade-offs are necessary with duplex systems. The major limitations when making Doppler measurements with duplex
systems are the footprint of the probe that limits access at the
optimal Doppler angle to some sites (left main coronary arteries), limited temporal resolution (needed for PWV and cardiac
timing), and limited resolution of higher velocities (needed for
estimating the pressure drop at a stenosis). For instance, we
have measured jet velocities as high as 4.5 m/s in TAC mice
and coronary velocities as high as 3.6 m/s in atherosclerotic
mice during hyperemia (21). From Eq. 1, the maximum velocity limit using a 30-MHz probe at a pulse repetition frequency
Drs. C. J. Hartley and A. K. Reddy collaborated extensively with Dr. S.
Madala of Indus Instruments on the design of the Doppler signal processing,
and much of the technology described in this review is available commercially
from Indus Instruments (Webster, TX). Indus also supplies procedure boards,
temperature controllers, and ECG amplifiers to VisualSonics for use with
mouse ultrasonic imaging systems.
AJP-Heart Circ Physiol • VOL
Summary
We have shown that high-resolution velocity and diameter
measurements can be made noninvasively and serially from
large and small peripheral vessels in mice. An advantage to
measuring blood velocity rather than volume flow is that the
values do not have to be normalized to body size (10, 21).
Because the lengths and time constants of the vascular systems
scale with the cardiac period (63, 64), reflected waves return at
similar times during the cardiac cycle, and the magnitudes and
waveforms of pressure and velocity in mice are almost indistinguishable from signals recorded from similar sites in humans and other species (42). Aortic PWV, percent arterial
diameter pulsation, ejection fraction, and CFR also have similar values, and many of the indexes (59) used to evaluate
cardiac and vascular function in mice are derived from human
indexes. Importantly, the use of familiar ultrasonic Doppler
methods and the similarity of velocity, pressure, and diameter
waveforms may permit much of what we learn about blood
flow and vascular mechanics in mice to be translated directly to
humans.
GRANTS
This work was supported in part by National Institutes of Health Grants
HL-22512, HL-42550, HL-57068, HL-52364, HL-42550, HL-42267, AG15568, HL-42313, HL-13870, HL-89792, and AG-13251.
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