IEEE TRANSACTIONS O N BIOMEDICAL ENGINEERING. VOL. 3X. NO. 8. AUGUST 1991

735

An Ultrasonic Method for Measuring Tissue Displacement: Technical Details and Validation for Measuring Myocardial Thickening Craig- J. Hartley, Senior Member, IEEE, Harvey Litowitz, Raphael S . Rabinovitz, Wei-Xi Zhu, Jacques E. Chelly, Lloyd H Michael, and Roberto Bolli

Abstract-We have developed a method for measuring myocardial thickening from a single ultrasonic transducer attached to the epicardium. Displacement of the underlying myocardial tissue is measured by following the phase of the echoes within a sample volume range-gated across the myocardium. The output is in the form of an analog signal. To verify the accuracy, resolution, and limitations of the system, we derived the equations relating the position of a reflector to the phase of its echo and compared the system output in vitro to a known input using a single moving target and a random distribution of scatterers, and in vivo to that of an ultrasonic transit-time dimension gauge. The results demonstrate that the 10 MHz system can accurately follow the motion of single or multiple targets with a resolution of 0.02 mm. In 25 dogs myocardial thickening measured with the displacement system compared favorably in both waveform and magnitude with thickening measured by the two-crystal transit-time method. Applications for the displacement method include: quantification of regional ventricular function in animal mqdels of cardiovascular diseases, measurement of endocardial to epicardial differences in the deformation of regional myocardium during the cardiac cycle, and evaluation of regional cardiac function in patients during and after corrective cardiac surgery.

INTRODUCTION HE QUANTITATIVE measurement of regional myocardial deformation during the cardiac cycle is important in studies of myocardial mechanics and in evaluating the effect of ischemia, infarction, and/or pharmacologic agents on myocardial function. Myocardial contraction (segment shortening) or the resultant wall thickening has been evaluated in the past by a number of techniques including implanted radiopaque beads [ 11 which are followed radiographically, or paired ultrasonic crystals [2] with the distance between them measured continuously by the transit-time principle. The time required for a pulse of sound emitted by one transducer to

T

Manuscript received March 4 , 1990: revised September 19, 1990. This work was supported in part by Grants HL-22512, HL-42550, HL-42267. HL-32161. HL-43151, and HL-26377 from the National Heart. Lung. and Blood Institute, and by a Grant-in-Aid from the American Heart Association, Texas Affiliate. The authors are with the Department of Medicine, Baylor College of Medicine, Houston, TX 77030. IEEE Log Number 9101 198.

be received by the other is related to their distance by the following equation: t, = d / c

(1)

where d is the distance between the transducers and c is the speed of sound in myocardium (1540 m/s). Although these types of transducers can be placed at any location and depth in the myocardium for cardiac dimension measurements, considerable skill is required in placing and aligning each pair. Despite the limitations, the ultrasonic transit-time method is commonly used to measure shortening or thickening in both acute [3], [4] and chronically instrumented animals [5], 161, and it has also been used both intraoperatively [7] and postoperatively [8] in a limited number of studies with patients. Various ultrasonic techniques have been used in the past to quantify the motion of heart structures using a single transducer applied to the skin surface [9] or directly to the heart [lo]. Because of commercial availability, the most widely used of these methods is M-mode echocardiography [ l 11. Briefly, pulses of ultrasound are emitted by a transducer placed over the area of interest and the retuming echoes from the various tissue interfaces are displayed. The return of each echo is related to the reflector distance by the following equation: t,, = 2 d / c

(2)

where d is the distance from the transducer to the reflecting interface and c is the speed of sound in tissue. The factor of 2 is included here because of the round-trip signal path. Since the speed of sound is relatively constant in most soft tissues at 1540 m / s (or 1.54 "Ips), the echo return time is proportional to the reflector distance. In M-mode devices, the echo amplitude is used to intensity modulate an oscilloscope trace which is swept across the display to show the motion of structures along the sound beam. This type of display has been clincially useful in displaying heart valve motion, chamber diameter, and wall thickness changes during the cardiac cycle. The resolution is limited to several wavelengths (0.5-1 .0 mm), and it is difficult to separate echoes from closely spaced

0018-929419110800-0735$01.OO 0 1991 IEEE

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IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING. VOL. 38. NO. 8, AUGUST 1991

structures such as adjacent layers within the myocardium. Also because the position of a structure must be traced by a stored or recorded image, motion is difficult to quantify in real time. To improve the resolution and provide a real-time output signal, others have built devices which track the leading edge of a well-defined echo [12], or which follow the zero-crossing of a poorly defined echo signal [ 131. These instruments have proven especially useful in displaying arterial wall motion noninvasively. Since the phase of an echo is related to its position, another approach to measuring tissue motion or displacement is to detect and process the phase rather than the amplitude of the returning echoes. One such device utilized continuous wave transmission and incorporated a phase-locked loop to track the echo from an arterial wall [14]. However, with each of these devices it has proven difficult to lock onto and remain locked to the desired echo, and our attempts to apply these techniques to the measurement of myocardial thickening have failed. To overcome the difficulties of tracking a poorly defined echo, we developed an open-loop approach in which phase is detected at a fixed distance from the transducer and followed over time to approximate the motion of the layers at that depth [15]. Similar techniques have been used with success by others to evaluate limited displacements such as vessel wall motion [16], [17] and tissue deformation [ 181 where motion is small relative to the size of the reflector. However, this method has not been applied previously for measuring large displacements of a deformable structure such as the myocardium. THEORYOF OPERATION The technique is based on principles of both pulse-echo and pulsed Doppler ultrasound [IS]. Briefly, using a transducer attached to the epicardium, myocardial thickening is measured by following the phase of the echoes from myocardial layers moving through a range-gated sample volume. The sample volume is approximately 3.6 mm3 and can be placed electronically at any depth within the myocardium. As the heart muscle thickens and thins during the cardiac cycle, muscle layers move back and forth through the sample volume. By detecting gnd following the phase of the echoes, the motion of the layers with respect to the epicardial transducer can be quantified. In the following we will show that the instantaneous phase of the returning echo is proportional to the distance of the reflector from the transducer. For simplicity, the equations will be shown for a single reflecting target in the sound beam as shown in Fig. 1. Using a tone burst mode of operation, the transmitted signal (S,) has the following form: S, = cos (ut), =

0,

for rzT

elsewhere

< t < nT

+ t,, (3)

where w is the angular frequency of the transmitted wave, T is the pulse repetition period which must be an integer

d

k

4

Sound) beam

St

S. Signals

Id

k

4

Fig. 1. Diagram showing the signal from one of the reflecting targets within the myocardium. The ultrasonic transducer is attached to the epicardium with the sound beam perpendicular to the surface. The time (t,,) between the transmitted signal (S,) and the received echo signal (S,) is proportional to the distance ( d ) from the transducer to the reflector. The actual signal from myocardium is a summation of signals from the thousands of reflecting sites distributed along the sound beam.

number of cycles of U , n is a positive integer, and t, is the duration of the transmitted burst. This tone burst is propagated from the transducer toward a target where part of it is reflected back toward the transducer which then acts as a receiver. The received echo signal (S,) has the form

S, = a cos {u(t for nT = 0,

tJ},

+ td < t

< nT

+ td + t,, (4)

elsewhere

where a is the amplitude of the received echo signal and td is the time delay between the beginning of the transmitted burst and the beginning of the received echo signal [see (2)]. Rearranging and susbstituting from (2) results in for r as in (4).

S, = a cos (ut - 2 w d / c ) , Since w

=

S,

(5)

2irc/X, ( 5 ) can be put in the form =

for t as in (4),

a cos (wt - +),

(6)

where X is the wavelength of the ultrasonic wave in the conducting medium (tissue), and where

4

=

47rd/X.

(7)

In this equation, 4 represents the phase (in radians) of the echo signal with respect to the transmitter signal and is directly proportional to the distance d from the transducer to the reflecting target. In the instrument, phase is sensed by a quadrature phase detector consisting of two analog multipliers. The reference inputs to the multipliers are cos ( a t )and sin (ut)and are derived from a master oscillator which runs continuously at angular frequency W . Multiplying each of these signals by the echo signal results in cos (ut) x a cos (wt - 4 ) = a(cos (2wr -

4)

+ cos (4)}/2

(8)

HARTLEY er a / . : ULTRASONIC DISPLACEMENT TECHNIQUE

737

and sin (ut) x a cos (ut - 4) =

a{sin (2ut

-

+)

+ sin ( + ) ) / 2

(9)

for t as in (4). If the above signals are low-pass filtered to remove the high frequency terms at 2ut, are multiplied by 2 to eliminate the 1/2, and are sampled only during the received interval [i.e., t as in (4)], the sampled, phase detected signals then become x = a cos

+

y

+

(10)

and =

a sin

of the target and in the magnitude and sign of its velocity if the sampling rate is not high enough. The above equations were derived for a single reflector in the sound beam. The actual received echo signal from myocardium, however, is not a single coherent burst, but is a summation of the individual echos from all the reflecting sites along the sound beam. These echoes interfere with each other in a random manner depending on their positions thus producing a complex and constantly changing signal. In addition, the displacement of individual targets during the cardiac cycle may exceed the sample volume length. The instrumentation must be able to process this signal and resolve the inherent ambiguities related to target motion.

(1 1) where x and y can be considered as components of a polar coordinate phase vector of length a and angle +. In general, the target will be moving with respect to the transducer which will generate a Doppler shift and cause a rotation of the phase vector. To see how the phase and Doppler shift frequency are related, consider a target moving at velocity v at an angle 8 with respect to the sound beam axis. Its distance from the transducer d is given by

INSTRUMENTATION A block diagram of the ultrasonic displacement measuring instrument is shown in Fig. 2. Timing for the instrument is controlled by the 10 MHz crystal oscillator shown to the left of center in Fig. 2. The 10 MHz frequency is divided by 2560 to produce a pulse repetition frequency (PRF) of 3.90625 kHz. A 0.4 ps pulse is used to gate 4 cycle bursts of the 10 MHz signal [see (3)] to the transmitter amplifier which drives the ultrasonic transd = v cos 8 dt = ut cos 8 + d,, (12) ducer. The transducer converts the electrical signals to acoustic tone bursts which are propagated into the tissue where do is the initial position of the target. If (12) is where they are reflected by structures along the sound substituted into (7), phase also becomes a function of time beam as shown in Fig. 1. The echoes returning to the transducer are converted back into electrical signals [see given by (4)] which are amplified and compared in phase to quad= (47rvt/X) cos 8 + constant. (13) rature signals {cos (ut) and sin (ut)) from the 10 MHz oscillator. The two-phase detector outputs [see (8) and Differentiating the phase yields an angular frequency (ad): (9)] are then sampled by a 0.2 ps range-gate pulse delayed wd = d+/dt = (47rv/h) COS 8. (14) by 2-50 ps from the transmit pulse. After sampling, the two signals are highpass filtered at 1 Hz to remove the dc Substituting u d = 27rfd and = c/f, (14) becomes components from stationary structures and low-pass filtered at l kHz to remove residual signals at the PRF. Exfd = (2fv/c) COS 8, (15) cept for the lower bandwidth, the sampled, filtered signals which is the well-known Doppler equation [ 191 with fd [see (10) and (1 l)] at this point resemble quadrature audio being the Doppler shift of the reflected wave. The Dopp- signals from a pulsed Doppler instrument for measuring ler shift can thus be obtained from the phase by differen- blood flow [20]. tiation. For a moving target the phase vector rotates with It is instructive to use the vector representation of the an angular frequency given by ud in a direction corre- quadrature signals as shown in the X-Y display in Fig. 2. sponding to the direction of target motion: counterclock- The radius a represents the amplitude of the echo from wise (CCW) for motion away from the transducer (ad- the target, and the phase of the echo represents the povancing phase) or clockwise (CW) for motion toward the sition of the target. The change in position (or displacetransducer (receding phase). ment) can be measured by noting the direction (CW or Because the sine and cosine functions are periodic at CCW) and counting the revolutions of the vector. Each 27r, (10) and (11) have ambiguities at intervals of d = revolution corresponds to reflector motion of 0.075 mm n X / 2 corresponding to target positions separated by 1 / 2 at 10 MHz. To improve the resolution of the instrument wavelength or 0.075 mm using 10 MHz ultrasound. In to 0.019 mm, revolutions are counted in 90" increments addition, the sampled x and y values change at discrete corresponding to axis crossings in the X-Y display of Fig. intervals separated by the pulse repetition period ( T ) . 2. Logic circuits in the up-down counter controller detect Thus, a given change in phase (A+) observed between positive and negative zero crossings of each quadrature two samples could be caused by an actual phase change signal, assign a direction based on the polarity of the zero of A$ + 2na where n is a positive or negative integer. crossing and the polarity of the other signal at the time, There are thus potential ambiguities in both the position and increment or decrement the 8-bit up-down counter.

i'

+

+

-

-

_ . _ _ ~ _ _ __

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IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING. VOL. 38. NO. 8. AUGUST 1991

Fig. 2. Block diagram of the ultrasonic displacement measurement system operating at 10 MHz showing the major steps in the signal processing. Equation numbers refer to those in the text.

Full-scale range for the 8-bit counter is - 4 . 8 mm. A digital-to-analog converter produces a voltage output which represents the change in position of echoes within the sample volume with a calibration of 2 V/mm. Since no filter is used on the displacement output, it is updated after each sample ( = 4 kHz) whenever the reflector has moved 0.019 mm. If the transducer is attached to the epicardium and the range-gate is set in the underlying myocardium, the output will display a waveform of the motion (thickening) within the sample volume. Because of transducer angulation and other sources of error, the up counts may not always precisely equal the down counts over the cardiac cycle, and the cumulative errors will cause the trace to wander up or down on the display or recorder. In addition, the starting position of the counter when the system is activated or when the transducer is connected is arbitrary. Two methods have been included to control these problems. The first is to inhibit the counter whenever it trys to exceed either zero in the negative direction or 255 in the positive direction. Counting resumes normally from 0 or 255 when the direction reverses. The first cardiac cycle after the system is activated will often be clipped, but subsequent cycles will usually be complete with minimal or no clipping. To allow small thickening signals to be enlarged on a recorder, it is necessary to further restrict the range of the counter to its lower end. A switchable upper limit of 6, 7, or 8 bits (63, 127, or 255 counts) keeps the trace within the range of a recorder set to 2.5, 5.0, or 10 V (1.25, 2.5, or 5.0 mm) full-scale. The second method is to reset the counter to a fixed number at a consistent point in the cardiac cycle such as the R-wave of the ECG. In this mode the counter is reset at the beginning of systole and any cumulative counting errors are taken out at each cardiac cycle. Signals available from the instrument include: displacement at 2 V/mm, analog range at 0.1 V/cm, and quadrature audio. In addition, oscilloscope monitor outputs are

provided from the transmit pulse generator, the range-gate pulse generator, and one of the phase detectors. Controls are a 2-40 mm range-gate potentiometer and a 6 , 7 , or 8 bit limit switch. Inputs are from a 10 MHz piezoelectric transducer attached to the tissue of interest, and a reset command from a triggered event such as the R-wave of the ECG. METHODS To test the accuracy, utility, and limitations of the displacement method for measuritig myocardial thickening, several studies were done in vitro using specially constructed simulators and in vivo using anesthetized and chronically instrumented dogs.

In Vitro Validation To evaluate the accuracy, frequency response, and linearity of the displacement method in measuring the motion of a discrete moving target, we constructed a test apparatus as shown in Fig. 3. A small wire (#40 AWG, 0.08 mm diameter) was attached to the center of the voice coil of a speaker which was suspended over a beaker of water. A 2 mm diameter 10 MHz ultrasonic transducer was placed at the bottom of the beaker directly under the end of the wire which was bent at 90". The height of the speaker could be adjusted vertically to position the wire anywhere alohg the sound beam from 0 to 4 cm from the face of the transducer. The speaker was driven by a triangle wave generator at an amplitude of 0 . 6 mm peak-topeak with a frequency of from 0.04 to 40 Hz producing target velocities from 0.05 to 50 mm/s. The amplitude and frequencies were chosen to keep the target echo within the sample volume throughout its motion and to span the range of velocities and range gate depths encountered in myocardial thickening measurements. The sample volume dimensions for the system were also measured using the apparatus shown in Fig. 3. To eval-

HARTLEY er

U/.:

ULTRASONlC DISPLACEMENT TECHNIQUE

739 Compressed

To signal generator

r---%q Transducer

Plexiglas

Transducers

Fig. 4. Apparatus for testing the performance of the displacement measurement system for a distribution of scatterers aimulated by a water-filled sponge. The transducers were mounted in plexiglass disks 5 cm in diameter above and below the sponge.

1 Fig. 3 . Apparatus for testing the frequency response, linearity, and resolution of the displacement measurement system for a discrete moving target. The speaker was mounted to a calibrated X - Y - 2 positioner so that the moving wire could be placed anywhere in the sound beam.

For epicardial thickening

For transmural

uate the lateral dimensions of the sample volume the wire was scanned across the beam at distances of 5, 10, and 20 mm from the transducer face with the speaker driven by a 10 Hz triangle wave. The amplitude was adjusted to give one complete loop of the phase vector as seen on an X-Y oscilloscope connected to the quadrature audio outputs of the instrument. The corresponding excursion is X/2 or 0.08 mm. An attenuator was placed in series with the transducer to prevent saturation of the RF amplifier and to keep the displayed loops circular at the highest signal level. The sample volume was centered on the wire and the detected audio power in decibels was plotted versus wire position in millimeters at each distance. To evaluate the axial dimensions of the sample volume, the wire was positioned in the center of the beam at distances of 5, 10, and 20 mm from the transducer face. At each position the audio power in decibels was plotted versus range gate delay in millimeters. To evaluate the performance of the system with a deformable distribution of scatterers whose individual displacements exceed the sample volume dimensions, we constructed a second test apparatus as shown in Fig. 4. The model consisted of a 27 mm thick water-filled sponge which could be compressed by known amounts from 1 to 5 mm to mimic the way the myocardium deforms during the cardiac cycle [21]. Transducers were placed both above and below the sponge so that each layer of the sponge could be sensed from either side. A trapezoidal waveform was used to compress the sponge at a frequency of 0.5 Hz and an amplitude of 2 mm. The gate depth was increased from 1.5 to 26 mm in 0.75 mm increments and measurements of peak-to-peak displacement were recorded from both the upper and lower transducers at each range depth.

In Vivo Validation The ability of the system to measure myocardial thickening in vivo was evaluated by comparing thickening measured by the single crystal displacement method to that measured by the two-crystal transit-time method in 25 open chest dogs [22]. One ultrasonic crystal was tunneled into the myocardium either half way (for epicardial thickening) or all the way to the endocardium (for trans-

Endocardium Transit-time crystal

Fig. 5 . Cross-sectional drawing of the myocardium of a dog showing the position of the displacement and transit-time transducers used in the in L'ivo validation.

mural thickening), and a second crystal was sutured to the epicardium over the first crystal for minimum transit-time as shown in Fig. 5. The epicardial crystal could be used simultaneously as both the transmitter of the two-crystal transit-time pair and as the transmitter-receiver for the displacement method. The range gate was set as close as possible to the mean depth of the deep crystal so that the measurements could be compared. To alter the regional ventricular function, several maneuvers were performed: occlusion of a coronary artery to produce ischemia, reperfusion of the coronary artery to produce partial recovery of function, infusion of isoproterenol to enhance contractility, and infusion of phenylephrine to depress contractility. At each level of function in each animal, thickening was measured by each method and compared. To facilitate comparisons between animals, systolic thickening was normalized to the end-diastolic thickness (transit-time method) or to the sample volume depth (displacement method) and expressed as percent thickening fraction (%TF).

RESULTS In Vitro Validation The results of the frequency response test at a target distance of 1 cm are shown in Fig. 6 . Each panel in the figure shows the speaker input below and the displacement output above. The frequency of the input triangle wave was 0.04, 1.O, and 40 Hz with an amplitude of 0.6 mm corresponding to target velocities of 0.05, 1.2, and 50 mm/s and to Doppler shift frequencies of 0.67, 16, and 667 Hz. The displacement waveform and the input

740

FREOUENCY TARGET VELOCITY

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 38. NO. 8, AUGUST 1991 .a4

.05

nZ

*M/SEC

1 ti2 1.2

40 50

HH/SEC

Hi

d S F '

DISTAIICE TO TARGET

2

4 cn

50

RANGE GATE DELAY

27

"SEC

1 cn

CM

0.5

13,s "SEC

"SEC

7

CM

TOUCHlNG

3

"SEC

I

INPUT

\

/

\

i

1 SEC. T I M E R

USEC

~

---~

Fig. 6 . Frequency response test of the displacement measurement system the apparatus shownin Fig. with a triangle wave input to the speaker at three different frequencies.

waveform are nearly identical over the 1000/1 range of target velocities demonstrating excellent fidelity, resolution, and stability at both high and low velocities. At the slower velocities the 0.19 mm steps can be seen in the output tracings. Fig. 7 shows the displacement output and speaker input for a 1 Hz triangle wave at target distances of 40, 20, 10, 5, and 0.0 mm. The waveforms are identical at each distance until the wire actually touches the crystal stopping its motion. Motion can thus be sensed anywhere along the sound beam with equal accuracy. Fig. 8 shows the results of the sample volume meaSignal strength SUrementS at ranges Of 5 , 10, and 20 ". in decibels (relative to the maximum signal at 5 mm) is plotted on the vertical axis reflector position in both axial and lateral dimensions at each distance. The position and the relative dimensions of the 2 mm diameter, unfocused transducer are also shown. The crossed, dotted lines on the 5 mm plots show the dimensions of the sample volume if the detection threshold is 20 dB down from the maximum signal. The results of the sponge compression test stimulating a distribution of scatterers are shown in Fig. 9. Each point represents the peak-to-peak displacement measured by the lower (solid points) or upper (open points) transducer at a specific range-gate depth while the sponge was compressed by a total of 2 mm. The curves were plotted in opposite directions to demonstrate that the upper and lower curves would overlap. Note that in this sponge the distribution of strain is nonuniform with the upper half compressing about twice as much as the lower half.

In vivo Vulidution The results of the in vivo validation are summarized in Figs. 10, 11, and 12. Fig. 10 is an X-Y plot of the quadrature signals from the myocardium of a dog during part of a cardiac cycle showing the locus of the phase vector. Two reversals in the direction of rotation can be seen in the plot. The slight scalloping of the loops is a result of the sampling process and the 1 kHz low-pass filter. Although the amplitude or radius varies considerably, the open area around the origin indicates that the revolutions and all axis crossings are clearly countable.

1 SEC. T I M E R

-

-

-~--- --

-

.

,..

-

.

~~

Fig, 7. Linearity test of the displacement measurement system at several target-to-transducer distances showing that the system can sense displacement from the face of the transducer out to at least 4 cm.

5"

Fig. 8. Curves showing the amplitude of the echo V S the position of the wire reflector with the range gate set to 5 , IO. and 20 mm using the apparatus shown in Fig. 3 . Both axial and lateral scans are shown at each distance. The amplitude scale is in decibels relative to the maximum signal. The relative size and position of the 2 mm diameter transducer is shown on the left

RANGE GATE DEPTH FROM TOP 20

I

I

10

I

I

I

MOOE)

1

m o m * ' 0 * e

MM

" 1

O t 0

MM

I 10 RANGE GATE D E P T H

el0

.o

27 MM SPONGE RUBBER COMPRESSED BY 2 MM (NON-TRACKING

0

MM

I

I 20

FROM

ROTTOM

t

1.3 5

2.0 0

Fig. 9. Plots of displacement versus range-gate depth for the upper and lower transducers as the 27 mm thick sponge shown in Fig. 4 was compressed by 2 mm. One curve was plotted from the lower left and the other from the upper right so that the measurement sites from each transducer would overlap and the curves from each could be compared.

Fig. 11 shows simultaneous waveforms of myocardial thickening measured by the single-crystal displacement method and the two-crystal transit-time method in a dog during control and ischemia created by occlusion of the left circumflex coronary artery. The waveforms obtained by the two methods are similar under each condition.

HARTLEY

PI

a l . : ULTRASONIC DISPLACEMENT TECHNIQUE

14 I

0

PHENYLEPHRINE

OPEN -TRANSMURAL SOLID

- EPICARDIAL 40

10%TF

-o:

y$

//

0

%TF

10 TRANSIT-T

20 IME

-I0-

-20-

Fig. 10. An X - Y oscilloscope display of the two phase detector outputs (0.25 s exposure) showing the locus of the phase vector from a displacement probe attached to the myocardium of a dog. Each complete revolution corresponds to displacement within the sample volume of 0.075 mm. Two reversals can be seen in the rotation indicating changes in the direction of motion (i.e., from thickening to thinning) within the sample volume. A digital counter is connected to count each transition of the vector from one quadrant (1, 2, 3, or 4) to the next. counting up for counterclockwise rotation (advancing phase) and down to clockwise rotation (receding phase). Thus, it counts up during systolic thickening, reverses, counts down during diastolic thinning, and finally retums to its initial value after one cardiac cycle. The counter output gives a waveform of displacement or thickening directly with each count corresponding to motion of 1 / 8 wavelength or 0.01875 mm. The sampling process causes some scalloping of the loops. and the stochastic nature of the scattering from the moving myocardium causes amplitude modulation of the signal which can be seen as variations in the diameter of the loops. As long as the loops encircle the origin so that the transitions between quadrants (1-2-3-4-I---. or4-3-2-1-4---) are proper, the phase counter will operate properly, and the instrument will accurately display the displacement within the sample volume.

CONTROL

LCX OCCLUDED

Fig. 12. Comparison of thickening fraction (%TF) measured by displacement and transit-time methods in a total of 25 dogs. Thickening was modified from control by ischemia and reperfusion and by administration of isoproterenol and phenylephrine. Both transmural (total wall) and epicardial (outer half) data are shown, and the number of dogs averaged for each data point is indicated. Some dogs were used for transmural data or epicardial data only, and not every intervention was performed in every animal 1221.

circles represent transmural measurements with the intramyocardial crystal and the sample volume placed in the subendocardium, and the solid symbols represent epicardial measurements with the intramyocardial crystal and the sample volume placed in the midwall. Transmural and epicardial measurements could not be made in all animals under all conditions because of difficulty in placing the intramyocardial crystal and loss of transit-time signals during the protocol. The number of animals averaged for each data point is indicated. Transmural T F varied from +20% during control and isoproterenol to -20% during ischemia. Epicardial TF varied from 1 1 % to - 14% under the same conditions. A 45" line of identity is shown for comparison of the two methods. A more detailed analysis of these data has been published elsewhere [22].

+

ECG

CORONARY LAD

"I\

12 THICKNESS TRANSIT-TIME

mm 10

Fig. 11. Waveforms from the displacement and transit-time measurements of myocardial thickening under control and ischemic conditions showing excellent agreement between the two methods.

Fig. 12 shows a comparison between thickening fraction measured by transit-time and displacement in a total of 25 dogs under several different conditions designed to alter thickening fraction from control values. The open

DISCUSSION The displacement method quantifies the motion or positional change of reflectors with respect to the transducer by following the phase of the echoes within the sample volume over time. The method involves several potential sources of error including: 1) ambiguities in the measured phase due to the sampling process and the periodicity of the sine and cosine functions, and 2) statistical uncertainties in the interference patterns of overlapping echoes due to the location of reflectors within the sample volume. Therefore, the accuracy and limitations of the method must be defined under various conditions. The ability of the method to quantify the motion of a discrete reflector over a wide range of target velocities and distances is demonstrated in Figs. 6 and 7. The magnitude and shape of the displacement output is indistinguishable from the speaker input voltage. In this setup, the motion of the wire target is small enough and the sensitivity of the signal processing is great enough that the echo is within the sample volume of the instrument (Fig. 8) throughout its 0.6 mm excursion. The small steps seen

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in the tracings of Fig. 7 show the changes in the least significant bit of the 8-bit counter which limits the resolution to h / 8 or 0.019 mm at 10 MHz. For myocardial thickening measurements this limitation results in an uncertainty of about 2 % when the total thickening is 1 .O mm. The frequency response of the measurement system is determined by the bandpass filters (nominally 1 Hz to 1 kHz) and the PRF (3.906 kHz), and was set to more than cover the range of velocities expected from myocardial thickening measurements. The range of Doppler frequencies covered by the data shown in Fig. 6 extends from 0.67 to 667 Hz. At higher frequencies, the response of the chart recorder was not adequate to display the signals without distortion. The displacement method is shown to be accurate over this range which covers velocities from 0.05 to 50 mm/s. Also, the ability of the system to respond to high accelerations is demonstrated by the sharp and undistorted peaks of the output triangle waves. Fig. 7 demonstrates that the system is equally sensitive and accurate at measuring displacement anywhere along the sound beam from the first detectable echo (in this case touching the transducer) out to at least 40 mm. Ventricular thickness varies from about 2 mm in rat hearts up to about 25 mm in hypertrophied canine and human hearts. The dimensions of the sample volume for a 2 mm diameter unfocussed air-backed transducer are shown in Fig. 8. The axial extent of the sample volume is determined by the instrumentation and is a function of the transmitter burst length, the range-gate pulse length, the bandwidth of the transducer, the bandwidth of the amplifiers in the signal path, and the threshold level of the signal processing. The lateral dimensions are determined by the shape of the sound beam, which is related to the size, shape, and focusing of the transducer and by the threshold level. In general, the axial shape is independent of distance from the transducer. The lateral shape, however, varies with distance and usually increases as the beam gets wider. Side lobes may also appear as seen at 20 mm in Fig. 8. A major determinant of sample volume size is the threshold level of the signal processing which affects both the lateral and axial dimensions. The threshold determines the minimum signal which will be processed and is normally set somewhere between the background noise level and the maximum received signal. At a given threshold setting, the sample volume size can be found from Fig. 8 by dropping down from the peak and drawing lines to intersect the curves. An example is shown in Fig. 8 for the 20 dB contour at 5 mm. For a circular transducer the volume should have circular symmetry and can be approximated by an ellipsoid of revolution with major and minor axes given by the crossed lines. The 20 dB sample volume at 5 mm for this transducer is about 3.6 mm'. For comparison, the 6 and 40 dB sample volumes at 5 mm are 0.71 and 18 mm3, respectively. For a uniformly reflecting material the echoes from the center of the sample volume will be much stronger and

generally will dominate the signal. However, an exceptionally strong reflector could overwhelm many weaker ones even though it is not centered in the sample volume. For example, we have found that when the sample volume is advanced beyond the endocardium and into the left ventricular chamber, the endocardial echoes can be detected when the sample volume is over 1 mm into the chamber because the echoes from blood are so much weaker than the echoes from the heart wall. The myocardium contains numerous closely spaced scattering sites which change in number and distribution constantly during myocardial contraction. The magnitude of thickening (1-5 mm for a 10 mm thick ventricle) is such that except for very close ranges, scatterers remain in the sample volume for only a short part of the cardiac cycle. This has a parallel with pulsed Doppler velocimetry where individual blood cells remain in the sample volume for only a short period of time, and yet their average velocity can be measured accurately [19], [20]. The displacement curves from the water-filled sponge shown in Fig. 9 demonstrate the ability of the system to measure the relative motion of a distribution of scatterers which deform in a nonuniform pattern similar to that predicted for myocardium. It has been calculated from thick-walled spherical and cylindrical models [2 11 and verified experimentally [2 I], [23], [24] that the subendocardial layers thicken more than the subepicardial layers by about 2 to 1 . This piece of sponge deformed in this manner when viewed from the bottom of the beaker where the lower transducer shows an increase in thickening with depth as indicated by the slope of the curve. Note that the upper transducer shows a curve of decreasing thickening with depth which when plotted upside down overlaps the curve from the lower transducer. Thickening fraction for the entire sponge is calculated by dividing the actual forced displacement (2 mm) by the initial thickness (27 mm) for thickening fraction (TF) of -2/27 = -7.4%. The minus sign is conventional since the sponge was compressed from its initial thickness. Thickening fraction can be estimated from the displacement output by dividing the measured displacement at maximum range (2 mm) by the range-gate depth (26 mm) [15] for a TF of -2/26 = - 7 . 7 % . Note that if the acutal thickening fraction were calculated using the compressed value of 25 mm as the initial thickness, the result would be +2/25 or + 8 % . The displacement system would give the same result as before ( + 2 / 2 6 = +7.7%) but with a positive sign. Thus, thickening fraction measured by the displacement system gives a value midway between values calculated using expanded or compressed values for initial thickness. In myocardium TF by displacement would approximate T F by transit-time normalized to the mean of end-systolic and end-diastolic thickness rather than to the conventional end-diastolic thickness [ 151. From a curve as shown in Fig. 9, thickening fraction can be measured in a layer-to-layer manner anywhere along the sound beam. Thickening fraction (one component of strain) can be defined more precisely as

HARTLEY et o l . : ULTRASONIC DISPLACEMENT TECHNIQUE

TF

=

AT/AR

=

dT/dR

143

(16)

where T is the measured thickening at range-gate depth R. Thus, the slope of the curve at any depth defines the local thickening fraction. In the case of the sponge shown in Fig. 9, TF increases from nearly 0 (flat slope) near the bottom to about 12.5% near the top. The overlapping curves show that both transducers measure the same distribution of compression. The scatter in the data points of Fig. 9 is a measure of the uncertainty of the measurement and is due to the nonuniformity in the size and distribution of scatterers along the sound beam(s). Highly reflective scatterers in the beam produce longer lasting echoes that can dominate smaller echoes when they are summed. This will produce discontinuities (longer and higher steps) in an otherwise smooth curve. Other characteristics could result from phase cancellations (echoes summed to nearly 0) or other counting errors due to the phase vector not encircling the origin (Fig. IO). Except for points near 14 mm (from the bottom and seen by both transducers), there is a monotonic increase in measured thickening with increasing depth. If many points are taken as shown here, a fairly smooth curve can be drawn, but if fewer points are taken errors can result. Because it is difficult to simulate completely myocardial echogenicity and contraction patterns in virro, we have done several validation studies in animals comparing the output of the single-crystal displacement system to that of a two-crystal sonomicrometry system [ 151, [22]. Some of these results have been published elsewhere [22] and will only be summarized here. Fig. 10 shows the locus of the phase vector from a transducer attached to the myocardium of a dog. Although the amplitude (radius) varies by a factor of 4 during this 0.2 s exposure, all the loops encircle the origin and are clearly countable. This shows that even though numerous scatterers pass in and out of the sample volume during the cardiac cycle, the resulting interference among echoes does not disrupt the phase enough to produce counting errors. In contrast to the transit-time method, the displacement method does not require that a specific layer be followed or that there be a well defined echo to track since displacement is measured at a fixed depth. The layer which is best represented by the displacement method is one which moves equally to either side of the sample volume during the cardiac cycle [15]. In the in vivo studies, therefore, the sample volume was set as close as possible to the mean depth of the deep crystal of the transit-time pair so that the measurements would be comparable. The waveforms and data illustrated in Figs. 11 and 12 show excellent agreement between the displacement method and the transit-time method in measuring myocardial thickening and thickening fraction in animals. The results are equally in agreement whether transmural (total wall) or epicardial (outer half) measurements are compared. The proximity of all the data points to the line of identity in Fig. 12 shows that there is no systematic over or under estimation

of thickening by the displacement method when compared to the transit-time method. In the above discussion we have ignored any changes in the speed of sound which may occur during cardiac contraction. Although we could find no reliable data from living, blood-perfused myocardium, we expect variations to be on the order of 1 % or less. Since the speed of sound affects the relationships between phase and displacement [see (7)] and between time and distance [see (2)] linearly, the maximum error in the determination of displacement or sample volume position would also be 1 % or less. However, in determining the dimensionless thickening fraction [see (16)], any changes in the speed of sound affect both numerator and denominator in the same way and thus cancel. APPLICATIONS The abiiity of this method to measure thickening versus depth is a unique capability which is not available with transit-time or echocardiographic methods which are limited to measurements at one or two specific depths [21], [25]. In addition, since no deep crystal or markers are required, the displacement method can measure thickening at more depths with less trauma and damage to the myocardium being measured [26]. It is also possible to utilize multiple range-gating to make measurements at several depths simultaneously from one transducer [27]. Fig. 13 shows systolic thickening versus depth curves for two different closed-chest, chronically instrumented, awake dogs. The upper curve shows uniform thickening fraction of 15.3% across the wall with the exception of the first 2 mm of the epicardium. The lower curve, however, shows an increase in thickening with depth ranging from 1 1 % near the epicardium to 42% near the endocardium with an average T F of 23 % . These curves are examples illustrating the range of values we have seen in conscious dogs. Others have shown using sonomicrometry and echocardiography with markers that thickening is nonuniform across the wall with the endocardial half of the wall contributing from 83 % [23] to 62% [21] of the total thickening. We have found that the shape of the thickening vs depth curves varies with position on the ventricle with thickening being more uniformly distributed near the equator of the ventricle (upper curve) than near the base or apex (lower curve) [28]. We have also shown that the shape of the curve changes with ischemia and during reperfusion with the endocardium suffering greater dysfunction and recovering more slowly than the epicardium [291. In addition to applying this method extensively in dogs, we have also used it to measure LV thickening in rats. Fig. 14 shows ECG, LV pressure, LV thickening, and aortic flow from an anesthetized rat. A 20 MHz version of the displacement instrument was used for this experiment to improve resolution to 0.01 mm, to allow the use of smaller (1 mm2) crystals, and to measure thickening at closer ranges. Even with the sensitivity increased to 0.625

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THICKENING ~~~

TF

=

15,34

MM

0 0

MM

5 10 RANGE ATE DEPTH

TFAvG

=

15

23%

1

"L TF

=

11%

0 0 MM 5 10 15 Fig. 13. Systolic thickening vs depth curves from two conscious, chronically instrumented dogs illustrating uniform (upper) and nonuniform (lower) thickening across the wall. Thickening fraction (TF) at each depth or layer is approximated by the slope of the curve.

E.C.G.

L.V. PRESSURE

L.V. WALL THICKENING 3 M M RANGE

AORT I C FLOW

VELOCITY

L 0.5 SECONDS d Fig. 14. Multiple cardiac signals from an anesthetized rat showing ECG, left ventricular (LV) pressure from a Millar Mikro-tip catheter, LV thickening from a 20 MHz displacement transducer, and aortic flow velocity from a 3.5 mm diam 20 MHz pulsed Doppler cuff. The lines to the right show the opening (0) and closing (C) of the mitral ( M ) and aortic ( A ) valves. Systolic thickening is about 0.5 mm at 3 mm deep for a calculated thickening fraction of 1 7 % . Heart rate is 230 beats per minute.

mm full-scale, the thickening signals are clean and reproducible. The 0.01 mm steps can be seen near the peaks of the thickening signal. We have recently begun using this method on patients during cardiac surgery to assess regional ventricular function, to detect infarcts and aneurysms, and to analyze dysfunctional areas. For this application the 10 MHz crystal

was mounted into a silicone rubber suction cup which could be moved rapidly and easily to different regions of the LV. Fig. 15 shows an example of signals from two patients, one of whom had a previous myocardial infarction. The upper signal is from a normally contracting region of the LV which thickens by 20% during systole denoted by the vertical bars. The middle signal is from the

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145

TF

=

+20%

L

THlCKENlNG

I SEC TIMER

I

I

---

TF

=

-12:

AT AIIEURYSM

ECG

TF

=

0%

THlCKtill NG lSCHEHIC

Fig. IS. Wall thickening signals from thre LV sites on two different patients during coronary artery bypass surgery using a suction cup probe. The beginning and end of systole are marked by the vertical bars in each panel. Systolic thickening fraction (TF) varies from -12% to + 2 0 % .

same patient but over an aneurysm caused by the infarct. This waveform, which thins by 12% during systole, resembles an inverted LV pressure signal and shows completely passive motion with no muscle activity and no viability. The lower signal is from a different patient whose ventricle had become ischemic during surgery. Although there is zero net systolic thickening, the early diastolic thickening indicates that the myocardium is still viable and may recover its function if it is reperfused. This method has also been adapted to measure the mechanical properties of soft tissues in the extremities of man [30], [31]. Using a forced oscillation at the skin surface, the distribution of strain in the underlying tissue can be measured and the elasticity of the various layers can be determined. Knowledge of the regional tissue properties of the residual limb of an amputee is essential in designing a prosthesis which fits properly.

CONCLUSIONS We have shown that the ultrasonic displacement method for measuring myocardial thickening is an accurate and reliable tool for quantifying regional ventricular function. The in vitro validations with the single target confirm theoretically predicted values for resolution (0.02 mm) and the ability to track low (0.05 mm/s) and fast (50 mm/s) moving targets. The sponge test shows the ability to measure nonuniform compression and to operate adequately with a deformable and random distribution of reflectors whose excursions exceed the dimensions of the sample volume. The in vivo studies in animals and in man demonstrate its accuracy in measuring actual ventricular thickening under experimental and clinical conditions.

The simplicity in applying the transducers to the heart makes the ultrasonic displacement method ideal for measuring regional ventricular function in patients during and after cardiac surgery. ACKNOWLEDGMENT The authors gratefully acknowledge the contributions of B. Washington in constructing the transducers used in these studies and J. Brooks in providing editorial assistance.

REFERENCES J. Heikkila, B. S . Tabakin, and P. G. Hugenholtz, “Quantification of function in normal and infarcted regions of the left ventricle,” Cardiovasc. Res., vol. 6, pp. 516-531, 1972. P. Theroux, D.Franklin, J. Ross, Jr., and W . S . Kemper, “Regional myocardial function during acute coronary artery occlusion and its modification by pharmacologic agents in the dog,” Cir. Res., vol. 35, pp. 896-908, 1974. K. P. Gallagher, T . Kumada, A . Battler, W . S . Kemper, and J. Ross, Jr., “Isoproterenol-induced myocardial dysfunction in dogs with coronary stenosis,” Am. J . Physiol. 242 (Heart Circ. Physiol.), vol. 1 1 , pp. H260-H267, 1982. K. P. Gallagher, T . Kumada, J . A . Koziol, M. D . McKown, W . S . Kemper, and J. Ross, Jr., “Significance of regional wall thickening abnormalities relative to transmural myocardial perfusion in anesthetized dogs,” Circularion., vol. 62, pp. 1266-1274, 1980. K. P. Gallagher, M. Matsuzaki, J . A . Koziol, W. S. Kemper, and J. Ross, Jr., “Regional myocardial perfusion and wall thickening during ischemia in conscious dogs,” Amer. J . Physiol., vol. 247, pp. H727-738, 1984. S. Sasayama. D. Franklin, J. Ross, Jr., W. S . Kemper, and D. McKown, “Dynamic changes in left ventricular wall thickness and their use in analyzing cardiac function in the conscious dog,” Amer. J . Cardiol.. vol. 38, pp. 870-879, 1976. J . D. Slack, P. F. Wigler, J. V . Zeck, H . G. Hanky, J . S . Cole, and T. A . Patrick, “ A regional ultrasonic technique to measure ventric-

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ular performance during coronary artery bypass,’‘ J . Curdio\usc.. Surg.. vol. 23, pp. 166-171. 1982. R. C. Hill, W. R. Chitwood, Jr.. J. D. Sink. J. L. Cox, and A. S . Wechsler, “Perioperative assessment of segmental left ventricular function in man,” Arch. Surg., vol. 1 15, pp. 609-614. 1980. C. R. Joyner and J . M. Reid, “Ultrasound cardiology in selection of patients for mitral valve surgery.” Ann. N . Y. Acad. S r i . , vol. 118, p. 512, 1965. D. A . Conetta. L. G . Christie, Jr.. W. W. Nichols, R. L. Feldman, C . J. Pepine, and C. R. Conti. ”Echocardiographic analysis of systolic and diastolic left ventricular wall motion during transient myocardial ischemia,” J . Clin. Ultrasourzd., vol. 9 , pp. 59-65. 1981. R. C . Waag and R. Gramiac, “Ultrasound Basics,” in Cardiac Ulrrasound, R. Gramiac and R. C. Waag, Eds. St. Louis, MO: Mosby, 1975, pp. 1-18. D. E. Hokanson, D. E. Strandness, Jr.. and C. W. Miller. “An echotracking system for recording arterial-wall motion.“ IEEE Trans. Sori. Ultrason.. vol. SU-17, pp. 130-132, 1970. D. E. Hokanson. D. J. Mozersky. D. S . Sumner, and D. E. Strandness, Jr., ”A phase-locked echo tracking system for recording arterial diameter changes in vivo.” J . Appl. Physiol.. vol. 32, pp. 728-733, 1972. D. W. Baker and V. E. Simmons, “Phase track techniques for detecting arterial blood vessel wall motion,” in Proc. 2lsr Ann. Conk Eng. Med. Biol., 1968. C. J . Hartley. L. A. Latson, L. H. Michael, C. L. Seidel. R. M. Lewis, and M. L. Entman, “Doppler measurement of myocardial thickening with a single epicardial transducer.” Amer. J . Physiol., vol. 245, pp. H1066-1072, 1983. F. D. McLeod, G. M. Graves. D. E. Dick, and C. W. Miller, “Doppler measurement of wall motion,” in Ultrasound in Medicine. Engineering Aspects. D. N. White and R. E. Brown. Eds. New York: Plenum, 1977, vol. 3B. pp. 1373-1374. A. P. G. Hoeks, C. J. Ruissen. P. Hick, and R. S . Reneman. “Transcutaneous detection of relative changes in artery diameter.” Ultrasound Med. Biol., vol. 11, pp. 51-59, 1985. L. S . Wilson and D. E. Robinson. “Ultrasonic measurement of small displacements and deformations of tissue.” Ultmsoti. h u g . . vol. 4 , pp. 71-82, 1982. D. W. Baker, “Pulsed ultrasonic Doppler blood How sensing,” lEEE Trans. Son. Ultrasound., vol. SU-17, pp. 170-185. 1970. C . J. Hartley and J. S . Cole, “An ultrasonic pulsed Doppler system for measuring blood flow in small vessels,” J . Appl. Physiol., vol. 27.pp.626-629, 1974. K. P. Gallagher. M. C. Sterling, M. Choy, C. A. Szpunar. R. A. Gerren, M. T . Botham, and J. H . Lemmer. “Dissociation between epicardial and transmural fupction during acute myocardial ischemia,” Circ,ulation, vol. 71. pp. 1279-1292. 1985. 1221 W. X. Zhu. M. L. Myers, C. J’.’Hartley, R. Roberta. and R. Bolli. “Validation of a single crystal for measuremcnt of transmural and eDicardial thickening,” Amer. J . Ph\siol., vol. 251. pp, H1045H1055, 1986 [23] H N. Sabbah. M Marzilli, and P D Stein. “The relative role ot subendocardium and subepicardium in left ventricular mechanics,” Amer J Physiol , vol. 240, pp H920-H926, 1981 (241 J H Myers, M C Stirling. M Choy, A J Buda, and K P Ghallagher, “Direct measurements of inner and outer wall thickening dynamics with epicardial echocardiography.” Crrculntrorz. vol. 74, pp 164-172, 1986 [ 2 5 ] N. G Pandian, D J Skorton. S . M. Collins, H L Falsetti, E R Burke, and R. E. Kerber, ”Heterogeneity of left ventricular seginental wall thickening and excursion in 2-dimen\ional echocardiogram$ of normal human subjects,” Amer. J . Curdiol.. vol. 51. pp. 16671673. 1983. 1261 C . I . Hartley, R. M . Lewis. L. H . Michael, and M. L. Entman, ”Dynamics of transmural myocardial thickening,” in Proc. 36th Ann. Conf. Eng. Med. Biol.. vol. 167, 1983. C. J . Hartley, H. Litowitz, W. X . Zhu, R. Bolli. and M . L. Entman, “Quantification of layer-by-layer myocardial thickening with an epicardial ultrasonic transducer,” Med. Sic$. Eng. Comp., vol. 23 Suppl. 2. pp. 1300-1301, 1985. C. J . Hartley, H . Litowitz, and R. M. Lewis, “Regional transmural distribution of myocardial thickening,” in Proc. 37th Ann. Con8 Eng. Med. Biol.,vol. 32.7, 1984. R. Bolli, B. S . Patel, C . J. Hartley, J. 1. Thornby, M. 0. Jeroudi, and R. Roberts, “Nonuniform transmural recovery of contractile

function in stunned myocardium,” Amer. J . Physiol.. vol. 257. pp. H375-H385, 1989. [30] T . A. Krouskop, D. R. Dougherty, and F. S . Vinson, “A pulsed Doppler ultrasonic system for making noninvasive measurements of the mechanical properties of soft tissue.” J . Rehab. Res. Develop. 24. pp. 1-8, 1987. [31] S . F . Levinson and V . L. Newhouse, “The phase response of ultrasound to vibration: A method of measuring tissue elasticity.” Ultrason. Inzug., vol. 10, p. 74, 1988.

Craig J . Hartley (S’64-M’66-SM’88) was born in Spokane, WA, on April 2 , 1944. He received the B.S.E.E. and Ph.D. degrees from the University of Washington. Seattle, in 1966 and 1970, respectively. From 1970 to 1973 he was a Postdoctoral Fellow and Research Engineer in the Bioengineering Laboratories at Rice University, Houston, TX. Since 1973 he has been with Baylor College of Medicine, Houston, TX and is currently a Professor in the Department of Medicine, Section of Cardiovascular Sciences. Since 1968 he has been active in the development of ultrasonic techniques to measure blood How and cardiovascular function. He is the Principle Investigator on several research grants and was the recipient of a Research Career Development Award from the National Institutes of Health. Dr. Hartley is a member of American Institute of Ultrasound in Medicine, The American Physiological Society, The Cardiovacular System Dynamics Society. Tau Beta Pi, and serves on the Central Research Review Committee of the America1 Heart Association, Texas Affiliate.

Harvey Litowitz received the B.A. degree in biology from Case Western Reserve University, Cleveland, OH, in 1974 and the M.S. degree in biomedical engineering from the University of Houston, Houston, TX, in 1986. He is currently a clinical research assistant in the Department of Cardiology at the Cleveland Clinic Foundation where his research interests include use of echo ultrasound to assess valve function.

Raphael S. Rabinovitz was born in Haifa, Israel, on January 18. 1954 He received the B Sc and M.Sc. degrees in mechanical engineering from the Technion. Haiti. Israel, in 1979 and 1981, and the Ph D degree in mechanical engineering from the University of Houston. TX, i n 1986 He is an Assistant Professor at Baylor College of Medicine. Houston, TX His major research interest includes implantable ultrasonic sensors for blood How and cardiac function

Wei-Xi Zhu was born in Shanghai, China, on November 10, 1953. He received the M.D. degree from Sichuan Medical College West China Union University in 1982. He completed two years of research fellowship in cardiology and three years of residency in internal medicine at Baylor College of Medicine in Houston, TX, between 1984 and 1989. Since then he has been a clinical cardiology fellow at the Mayo Clinic. Rochester, MN.

HARTLEY e /

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ULTRASONIC DISPLACFMFNT TECHNIOUF

Jacques E. Chelly was born i n Paris, France. on September 3, 1949. He received the B.S. degree in biology from Mont-Rouge College, Paris. in 1970, the M.D. degree from Necker Enfants-Malades Medical School, Paris, in 1976. the M.S. in pharmacology from Lariboisikre-St. Louis Mcdical School. Paris, in 1979, and the Ph.D. degree from the University of Houston. Houston. TX, in 1985. He is Professor of Anesthesiology and Pharmacology, Director of the Division of Clinical Pharmacology, and Director of Clinical Research, Department of Anesthesiology, at the University of Texas Medical School at Houston. His current areas of research are cardiovascular anesthesiology and pharmacology.

Lloyd H. Michael received d n dssociate of a m degree from Keystone Jr College, La Plume, PA, in 1962. the B S degree i n Biology from Moravian College, Bethlehem, PA. 1964. the M S degree i n animal physiology from Kent State Uni vervty, Kent, OH, i n 1966, and the Ph D degree in physiology from the Univer\ity of Ottawa Medical School, Ottawa, Canada He completed poutdoctordl fellowships i n inyocardial biology at Baylor College ot Medicine, Houston, TX. and clinical immunology (Columbia College ot Phy sicians and Surgeons)

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He is Professor of Medicine. Molecular Physiology and Biophysics at Baylor College of Medicine, Houston. TX. Current research areas include: early myocardial ischemia and animal models to investigate the role of inflammation. leukocytes, and complement in myocardial darnage: the mechanics of cardiac muscle contraction.

Roberto Bolli was born in Perugia, Italy. on March 9, 1952. He received the M.D. degree from the University of Perugia in 1976. After two years of training in Internal Medicine at the University of Perugia. he began a two and one half-year Research Fellowship at the Cardiology Branch. NHLBI, NIH, from 1978 to 1980. He then completed a Clinical Fellowship in Cardiology at Baylor College of Medicine between 1981 and 1983, following which he joined the Faculty in the Cardiology Section at Baylor College of Medicine. He is now Associate Professor of Medicine, Director of the Experimental Animal Laboratory and Director of the Coronary Care Unit at the Veterans Administration Medical Center. His research interests include protection of ischemic myocardium, mechanisms of ischemia and reperfusion injury, postischemic myocardial dysfunction (stunned myocardium) and measurement of regional myocardial function with the ultrasonic technique developed by Dr. Hartley.

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