Integrated catheter for 3-D intracardiac

ieee transactions on ultrasonics, ferroelectrics, and frequency control, vol. 51, no. 7, july 2004
799
Integrated Catheter for 3-D Intracardiac
Echocardiography and Ultrasound Ablation
Kenneth L. Gentry and Stephen W. Smith, Member, IEEE
Abstract—A catheter device with integrated ultrasound
imaging array and ultrasound ablation transducer is introduced. This device has been designed for use in interventional cardiac procedures in which the cardiac anatomy is
first imaged using real-time three-dimensional (3-D) ultrasound, then ablated to treat arrhythmias. The imaging array includes 112 elements operating at 5.4 MHz arranged
in a 2-D matrix. Individual elements have a bandwidth of
21% and an insertion loss of 80 dB. The array has an azimuth resolution of 12 and an elevation resolution of 8.7 .
The ablation transducer is a concentric piezoelectric transducer PZT-4 ring (outside diameter (O.D.), 4.5 mm, inside
diameter (I.D.), 3.1 mm) operating at 10 MHz that surrounds the imaging array. It can produce a spatial-peak,
temporal-average intensity up to 16 W/cm2 . The entire device fits into a 9 Fr lumen with a 14 Fr tip to accommodate
the ablation ring. With this device we have imaged, in realtime 3-D, a variety of targets including wire phantoms, fixed
sheep hearts, and fresh bovine tissue. The ablation ring has
been used to heat tissue-mimicking rubber 14 C, as well as
create lesions in fresh bovine tissue.
I. Introduction
trial fibrillation (AF) is the most common cardiac
arrhythmia. This disease affects up to 2 million Americans with 160,000 new cases every year, and it is associated with both increased morbidity and mortality [1]–[3].
Radiofrequency (RF) catheter ablation is the most widely
used treatment for AF. It seeks to block conduction of arrhythmic electrophysiological waves from the sites of AF
initiation and maintenance (usually in or near the pulmonary veins) to other parts of the atria by creating transmural lesions in the left atrial wall [4]. A catheter with
multiple electrodes 4–8 mm in length is positioned in contact with the left atrium. An RF generator produces up
to 50 W of power at 300–1000 kHz, which travels from the
electrodes through the body and into a dispersive ground
pad on the patient’s skin [5]. Resistance in the interface
between the catheter tip and the heart tissue causes heating; keeping the electrode tip at 50–65◦C for 30–60 seconds
creates lesions [5]–[7]. Catheter placement within the heart
usually is performed under fluoroscopic guidance. Long
procedure times associated with RF catheter ablation increase the risks associated with long-term exposure from
ionizing radiation to the patient as well as medical person-
A
Manuscript received June 20, 2003; accepted March 17, 2004. This
research was supported by the NIH (grants HL64962 and HL72840)
and the NSF (grant DMR9973801).
The authors are with the Department of Biomedical Engineering,
Duke University, Durham, NC (e-mail: [email protected]).
nel. Additionally, RF ablation is associated with stenosis
of the pulmonary vein.
In an attempt to reduce fluoroscopy exposure, researchers have used intracardiac echocardiography (ICE)
to visualize heart anatomy during RF catheter ablation procedures. A 9 Fr, 9-MHz, rotating single-element
catheter (EP Technologies, Boston Scientific Corp., San
Jose, CA) produces a radial scan useful for imaging a cross
section of the pulmonary veins [7]. (The outside diameter
of a lumen is measured in French, or Fr, 1 Fr ≈ 0.33 mm).
After the ablation is performed, ultrasound is used to monitor for stenosis of the vessel. A 10 Fr, 5.5–10 MHz, phased
array ultrasound catheter (Acuson Corp., Mountain View,
CA) also has been used to guide RF ablation [8]. In this
case the ultrasound catheter is kept in the right atrium
where it can image both atria and the RF catheter. Both
of these commercial devices are limited to planar B-scan
images.
In real-time, 3-D (RT3D) imaging, a 2-D array transducer is used to steer and focus the ultrasound beam over
a pyramidal volume [9], [10]. The Model 1 commercial ultrasound system (Volumetrics Medical Imaging, Durham,
NC) can acquire this pyramidal volume and display up to
five simultaneous image planes from within this volume, as
well as RT3D rendered images. Clinical and animal evaluations have shown potential advantages over conventional
2-D scanners for measurement of cardiac function in terms
of ventricular volumes [11], peak left ventricular flow velocities [12], and perfusion [13]. In intracardiac RT3D imaging, the 2-D array is built into a catheter that then is
guided into the heart. Several RT3D catheters have been
described by our laboratory; a 5 MHz, 12 Fr side-viewing
device [14], [15], a 7 MHz, 9 Fr side-viewing device [16]; and
most recently, 5 MHz, 14 and 22 Fr forward-viewing devices with integrated working lumens [17]. These catheters
have visualized heart anatomy, including pulmonary vein
ostia, and been used to guide a number of surgical procedures in vivo in sheep, including RF ablation.
Catheter-delivered ultrasound also has been investigated for use in performing the actual ablations necessary to treat AF. The amount of energy transferred from
the acoustic wave to the tissue is directly proportional to
both the intensity of the wave and the absorption coefficient of the tissue [18]. Thus, if the ultrasound ablation
transducer transmits into a medium with low absorption
(e.g., water, blood), unlike RF ablation the catheter tip
need not be in direct contact with the myocardium. Zimmer et al. [19] created transmural lesions in canine cardiac tissue using transducers producing spatial-average,
c 2004 IEEE
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ieee transactions on ultrasonics, ferroelectrics, and frequency control, vol. 51, no. 7, july 2004
temporal-average intensities (ISATA ) of 5–15 W/cm2 for
60 seconds. A commercial ablation transducer (Atrionix,
Inc., Sunnyvale, CA) was reported to have created circumferential lesions in human pulmonary vein ostia after
a 2-minute ablation procedure [20]. These researchers also
have reported a lack of pulmonary vein stenosis after ultrasound ablation. In animal studies, Azegami et al. [21]
used ICE to assess pulmonary vein size before and after
ultrasound ablation. The maximum power applied to the
ablation transducer was 40 W for 30–240 seconds. Most
recently, Meininger et al. [22] reported successful circumferential lesion creation outside canine pulmonary veins
using a circular transducer (Transurgical, Setauket, NY)
emitting 40 W acoustic power for 30-120 seconds.
This paper describes a combined catheter capable of
both RT3D ICE and ultrasound ablation [Fig. 1(a)]. Such
a device could greatly simplify cardiac procedures in which
the heart anatomy must first be imaged then ablated,
such as to treat AF. A clinical device must image both
the cardiac chambers and pulmonary vein ostia and create transmural lesions at targeted locations. Previously, we
described a combined 12 Fr, side-viewing catheter with a
5 MHz imaging array, adjacent to a 10 MHz ablation element, and five integrated electrocardiogram (ECG) electrodes [Fig. 1(b)] [23]. This device imaged wire phantoms and ex vivo bovine tissue. The ablation element produced a spatial-peak, temporal-average intensity (ISPTA )
of 30 W/cm2 . At this intensity, a 56-mm3 transmural lesion was created in the 7-mm thick bovine tissue. Because
of the ablation element’s placement adjacent to the imaging array, direct visualization of ablation sites was difficult. Herein, we describe our efforts to improve on this
device by increasing imaging resolution and designing the
ablation beam and imaging volume to be coincident. In
Section II we describe design and fabrication of a 14 Fr,
forward-viewing catheter with a 5-MHz array for RT3D
ICE and a 10-MHz ring transducer that fits around the
imaging array [Fig. 1(c)]. In Section III, imaging of wire
phantoms and an ex vivo sheep heart and ablation studies
on phantom materials and ex vivo bovine tissue are described. In Section IV, our results are analyzed and future
work is described.
(a)
(b)
II. Methods
A. Transducer Design and Fabrication
The imaging part of our combined transducer is based
on a proven design [17]. As shown in Fig. 1(c), the
imaging array (no shading) has 112 elements on a 150µm pitch arranged in a 10 by 14 matrix. Each element operates at 5 MHz. In an earlier paper [24],
the two-way acoustic beam from this array was simulated using Field II software (available for download at
http://www.es.oersted.dtu.dk/staff/jaj/field) and found
to have a −6 dB width of 12◦ in azimuth and 8.7◦ in elevation (Fig. 2). This transducer was built with methods
(c)
Fig. 1. (a) Schematic of proposed device operation showing RT3D
imaging and ablation from a single catheter. (b) Photograph of previously made prototype device showing ablation transducer (A), imaging transducer array (I), and electrocardiogram electrodes (E). Top,
face view; bottom, side view. (c) Schematic of proposed imaging array (no shading), grounded elements (dark gray), and ring transducer
(light gray).
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Fig. 2. Simulated imaging array beamplots in the elevation direction, left, and azimuth direction.
Fig. 3. Simulated ablation ring beamplot.
described previously [25] using PZT-5H on a multilayer
flexible interconnect circuit (MicroConnex, Snoqualmie,
WA) backed by low impedance polymer (Z = 3.8 MRayls,
α = 1.1 dB/mm/MHz). Sixteen ribbon cables with 14 conductors per ribbon [seven signal channels alternated with
grounded conductors (Microflat, W. L. Gore & Associates,
Pleinfeld, Germany)] were fed through a 9 Fr catheter lumen to electrically connect the solder pads on the flex circuit to the system cable of our 3-D scanner.
The size of the ablation transducer [light gray in
Fig. 1(c)] was constrained by the design of the imaging
array. A ring of grounded elements (dark gray) surrounds
the imaging array, thus the minimum I.D. of the ablation transducer is 3 mm; our ring has a 3.1-mm I.D. The
O.D. was chosen to be 4.5 mm, both to keep the overall
diameter of the device small (14 Fr) and to ensure an electrical impedance well matched to our driving hardware.
The transmitted acoustic beam of the ablation transducer
was simulated using Field II software and found to have a
−6 dB width of 2.0◦ and a −20 dB width of 22◦ (Fig. 3). A
10-MHz, PZT-4 ring (VP-A40, Boston Piezo-Optics, Inc.,
Bellingham, MA) was centered on the imaging array and
Fig. 4. Face of completed integrated imaging and ablation transducer.
epoxied to the flex circuit. The PZT-4 was chosen for the
ablation transducer because its higher Curie temperature
and lower dielectric losses make it more suitable for highpower applications in comparison to PZT-5H. Additionally, PZT-4 has a higher mechanical Q and fewer mechanical losses than PZT-5H. The ring was mounted on the
imaging transducer so approximately half of it was backed
by the multilayer, flexible interconnect circuit and backing
epoxy with the remainder air backed. A miniature coaxial
cable was run through the catheter lumen and soldered
to the ring; the proximal end was terminated to a British
Naval Connector (BNC) connector. Heat-shrink tubing approximately 1 cm long was used to transition from the
smaller catheter lumen to the ablation ring and complete
the device (Fig. 4). No matching layers were used for either the imaging array or the ablation transducer in this
prototype device.
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ieee transactions on ultrasonics, ferroelectrics, and frequency control, vol. 51, no. 7, july 2004
B. Transducer Characterization
The impulse response of the imaging array was measured by transmitting from one element with a Model
9073PR pulser/receiver (Panametrics, Waltham, MA) and
receiving the reflection from an aluminum block with a
nearby element. The power spectrum was measured using
a Panametrics Model 5605A stepless gate and a Model
3588A spectrum analyzer (Hewlett-Packard, Palo Alto,
CA). The 50 Ω insertion loss of individual elements operating in both transmit and receive also was measured by
sending a pulse and receiving the echo from an aluminum
block 5 mm in front of the transducer. The imaging resolution of the system/array was measured by using the 3-D
scanner to image a 400-µm diameter sapphire sphere embedded in an acoustically transparent gel pad 4.7 cm from
the transducer. The brightness of the reflection from the
sphere was captured in a C scan parallel to the transducer
face. A cross section of this image, which approximates the
system beamplot of the array, was compared to a Field II
simulation of the array in which dead elements, element
bandwidth, and distance between transducer and target
were taken into account. The parallel processing performed
by the 3-D scanner widens the beam in comparison to the
simulation beam [9], [10]. This effect was not included in
the simulation.
The electrical impedance of the ablation ring transducer was measured with a Hewlett-Packard 4194A
impedance analyzer. In subsequent experiments, the ablation transducer was excited by a signal source (HewlettPackard 8165A) producing small-amplitude (<0.5 V),
high-frequency bursts with arbitrary cycle length and
pulse repetition frequency (PRF). These bursts were amplified by an RF power amplifier with 50 dB of gain
(Model 525 LA, ENI, Rochester, NY). Between the amplifier and the transducer, a power meter (Model NRT, Rohde & Schwartz, Munich, Germany) measured the power
available from the amplifier (forward power) as well as
the power reflected from the imperfectly matched transducer (reverse power). The width of the ablation beam
was measured by scanning a Model 804 calibrated membrane PVDF hydrophone (0.6-mm spot diameter, Sonic
Technologies, Hatboro, PA) in a plane parallel to the transducer face and recording the voltages as measured with a
Model 744A digitizing oscilloscope (Tektronix, Wilsonville,
OR). After finding the peak voltage at a range of 2 cm,
the magnitude of the voltage along four radii was measured and averaged. The same setup was used to find the
ISPTA produced by the ablation transducer according to
the procedures outlined by the Center for Devices and Radiological Health of the FDA [26]. Intensity was measured
first as a function of frequency in order to find the optimum center frequency of the device at which the least
amount of power is reflected from the transducer. Then
intensity was measured as a function of the number of cycles in the excitation burst. The PRF was 10.3 kHz for all
of the above experiments.
Fig. 5. Schematic of experimental setup for testing the ablation transducer. The catheter is shown at the top suspended in water. Five
thermocouples (black dots, not to scale) are shown embedded in the
attenuating rubber at a depth of 5 mm. Absorbing rubber is placed
at the bottom of the water tank to prevent any reflections.
C. Measurements
Our device was used to image the AIUM 100-mm wire
phantom with the Volumetrics scanner displaying both 65◦
and 90◦ pyramidal fields of view. Next, the ability of the
device to visualize heart anatomy was tested by imaging a
fixed sheep heart with a 90◦ field of view.
The ability of the ablation transducer to create lesions
was first tested in a tissue phantom material (polyvinyl
alcohol cryogel and graphite [27]) with an attenuation coefficient of approximately 0.5 dB/cm/MHz. Five type T
thermocouples (0.127-mm diameter wire, 5TC-TT-T-3636, Omega Engineering, Stamford, CT) were embedded in
the 2-cm thick, tissue-mimicking rubber as shown in Fig. 5.
The ablation transducer was suspended in water so that it
was above the phantom. A digital multimeter/data acquisition system (Model 2700 with 7708 multiplexing module,
Keithley Instruments, Cleveland, OH) under the control of
a laptop computer running LabVIEW software (National
Instruments, Austin, TX) was used to collect temperature
data from all five thermocouples once the signal source
had been enabled. After a selected amount of time passed,
the thermocouples stopped measuring and the temperature data was passed from the multimeter to the computer for display. For these experiments, the signal source
was set to produce 500 cycle bursts with an amplitude of
0.2 V and a frequency of 10 MHz with a PRF of 10.2 kHz.
The total duty cycle of the excitation pulse was 51% and
the transducer was powered for 60 seconds. The ablation
transducer was first positioned so that it was nearly touching the phantom, then a second procedure was run with it
positioned 5 mm above the rubber.
The ability of the ablation transducer to produce lesions in ex vivo beef muscle was tested. A hole 12 mm
in diameter was bored in a piece of beef 12–13-mm thick
to approximate an ostium in the wall of the myocardium.
For these first experiments, we approximated a degassed
system by using deionized water and massaging the tissue
to remove air bubbles. The catheter was suspended above
the beef and the Volumetrics scanner was used to image
the model. Then, the device was lowered to the surface
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Fig. 6. (a) Impulse response of typical imaging array element. (b) Power spectrum of typical imaging array element.
Fig. 7. (a) Comparison of experimental and simulated imaging array beam profiles in elevation direction. (b) Comparison of experimental
and simulated imaging array beam profiles in azimuth direction.
of the beef, and the ablation transducer was used to create a lesion near the lumen. The signal source was set to
produce 300 cycle bursts with an amplitude of 0.2 V and
a frequency of 10 MHz with a PRF of 17.6 kHz. The total duty cycle of the excitation pulse was 53%, and the
ablation lasted for 2 minutes. In total, three lesions were
created. The catheter was raised to its original position,
and the model was again imaged.
III. Results
A. Transducer Characterization
The typical impulse response of an imaging array element is shown in Fig. 6(a). Fig. 6(b) shows the corresponding power spectrum. The elements had a center frequency
of 5.4 MHz and a −6 dB bandwidth of 21%. Without accounting for diffraction losses, the average of 50 Ω insertion
loss measurements of six pairs of elements was 80 dB. The
azimuth and elevation beam profiles at a depth of 4.7 cm
were measured using the C scan image of the sapphire
bead and compared to beam profiles simulated with Field
II (Fig. 7). In the azimuth direction, a −6 dB width of
11.5 mm was measured as compared to a predicted width
of 8.4 mm. In the elevation direction, a −6 dB width of
8.7 mm was measured as compared to a predicted width
of 5.8 mm.
The ablation ring transducer had an impedance of 89 Ω
at the parallel resonance (10.1 MHz) of the thickness mode.
This was a good match to the 50 Ω output impedance of
the power amplifier, so no electrical matching was used.
The resonance at 2.55 MHz is caused by vibration in the
width of the ring. In Fig. 8, the measured and simulated
beam profiles of the ablation ring at a distance of 2 cm
are compared. The measured beam had a −6 dB width of
1.1 mm compared to 0.83 mm for the simulated beam. The
optimum center frequency of the ablation transducer was
10.0 MHz. At this frequency the amount of power reflected
from the transducer varied from 14% at low duty cycle
(1%, or 10 cycles per burst) to only 6.8% at high duty cycle (52%, or 500 cycles per burst). The low reflected power
may indicate an even better match to 50 Ω during highpower operation. The maximum ISPTA measured in this
study was 16.1 W/cm2 at 500 cycles per burst. At that
intensity, the amplifier provided 5.9 W to the transducer
according to the power meter. In Fig. 9, both ISPTA and
forward power are plotted as a function of burst length.
Forward power increases linearly with burst length, as expected, but ISPTA is linear only up to 100 cycles per burst.
B. Tissue Phantoms
Fig. 10 shows images of the AIUM wire phantom taken
with the catheter 2-D array. Fig. 10(a) shows a schematic
of the phantom with the separation of the wires. Fig. 10(b)
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Fig. 8. Comparison of experimental and simulated ablation beam
profiles.
and 5 only recorded a temperature rise of a few degrees.
Fig. 12(b) shows the temperature rise achieved when the
ablation transducer was positioned 5 mm above the absorbing rubber. In this case, thermocouple 3 was heated
only 8◦ C, thermocouples 2 and 4 were heated 3◦ C, and
thermocouples 1 and 5 were heated only a few degrees
each. For both of these experiments, the amplifier was providing 5.4 W of forward power with 5% of that reflected.
In Fig. 13, the results of our combined imaging and
ablation experiment are shown. Figs. 13(a) and (b) are a
B-scan and C-scan, respectively, of the hole in the beef.
Fig. 13(c) shows the results of one of the ablation experiments. The three lesions had an average depth of 1.3 mm
(standard deviation = 0.6 mm), an average diameter on
the surface of 4.1 mm (standard deviation = 0.9 mm),
and an average volume of 15.8 mm3 (standard deviation = 1.4 mm3 ). No images were viewable during the ablation procedure due to interference from the ablation signal, and the postablation images looked no different from
those in Figs. 13(a) and (b). During the ablation procedure, the amplifier provided 5.8 W of forward power with
14% reflected.
IV. Discussion
Fig. 9. Ispta and the electrical power available to the transducer are
shown as a function of the number of cycles per burst.
shows an axial view of the wires. In Fig. 10(c) a lateral
view in the elevation direction is shown. The two wires
1 mm apart are not resolved in this view. Fig. 10(d) shows
a RT3D rendered view of the wires. In Fig. 11, images of
the fixed sheep heart are shown. The catheter was positioned in the right ventricle (RV) of the heart and was
able to clearly image the RV, the left ventricle (LV) and
the septum separating the two chambers of the heart.
With the ablation transducer just above the tissuemimicking rubber, it was able to raise the temperature
of the tissue-mimicking rubber by 14◦ C as shown in
Fig. 12(a). The numbering of the traces refers to the numbering of the thermocouples in Fig. 5. Fig. 12(a) shows
that thermocouple 3, directly under the ablation transducer, reached a steady-state temperature change of 14◦ C
in 45 seconds. Thermocouples 2 and 4 reached a temperature change of approximately 7◦ C, and thermocouples 1
We have developed a prototype catheter with integrated
RT3D ICE and ultrasound ablation that may be useful
in interventional cardiac procedures in which the cardiac
anatomy is first imaged then ablated to treat arrhythmias.
We have presented images made by this device of a wire
phantom, a fixed sheep heart, and fresh bovine tissue. The
imaging array has resolution sufficient for the visualization of important cardiac anatomy, such as the cardiac
chambers and septum. The ablation transducer has heated
tissue-mimicking rubber by 14◦ C. We have shown that this
heating is adequate for creating lesions in bovine muscle.
The current prototype design leaves room for improvements in many areas. Imaging array resolution and sensitivity can be improved by increasing imaging frequency
[16] and by using matching layers, respectively. In order
to more directly monitor the ablation procedure, an ECG
electrode and thermocouple could be introduced in the
face of the transducer. The electrode would monitor for
the cessation of arrhythmia that should accompany cell
death, and the thermocouple would monitor temperature
at the transducer face. Research has shown that temperatures on the tissue surface in excess of 60◦ C enhance the
success of ultrasound ablation [28]. This thermocouple also
would be used to ensure that the temperature at the transducer surface never exceeds 100◦ C, the point at which the
blood boils. A lower operating temperature also extends
the life of the piezoelectric material in the ablation ring
transducer. Furthermore, in an attempt to simplify the
construction of this prototype catheter, we have not used
any acoustic matching layers or electrical impedance tuning of the ablation ring, which could increase the acoustic
output of the device.
gentry and smith: new catheter capable of rt3d ice and ultrasound ablation
805
(a)
(b)
(c)
(d)
Fig. 10. (a) Schematic of wires in phantom. (b) Axial view of wires. (c) Latral view of wires in elevation direction. (d) 3-D rendered view of
the wires.
Fig. 11. Images of fixed sheep heart. On the left, a long axis view of the left ventricle (LV) is shown, along with part of the right ventricle
(RV) and the septum. On the right, a simultaneous orthogonal scan is shown.
V. Summary
One main limitation of the device is its size. The 9 Fr
lumen is small enough for clinical use. By placing the ablation transducer around the imaging array, we have ensured that the ablation site is in the device’s field of view,
but we have made the tip too large (14 Fr) for routine
use in the clinic. One possible solution to the problem of
device size would be a single 2-D array for both imaging
and ablation. In imaging mode, the new array could be
slightly larger than the array presented here, improving
imaging resolution. Higher ablation ISPTA also would be
possible because of the increased surface area devoted to
ablation and focusing of the ablation beam. An additional
advantage is that the ablation beam could be steered in
2-D to create lesions of nearly any size or shape. Of course,
such a device would use the same piezoelectric material for
both imaging and ablation. If conventional PZT ceramics
(like PZT-5H and PZT-4 presented here) were used either
imaging or array performance will be compromised.
Acknowledgments
The authors thank Edward D. Light and Warren Lee
for their assistance in this project.
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Fig. 12. Temperature rise versus temperature in the five thermocouples shown in Fig. 5 and with the ablation transducer (a) nearly touching
the absorbing rubber, and (b) 5 mm above the absorbing rugger.
(a)
(b)
(c)
Fig. 13. (a) Image of beef tissue with 12-mm diameter hole. (b) C-scan of hole in beef tissue. The arrows in (a) indicate the plane of this
C-scan. (c) Photograph of lesion to the left of the hole.
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Kenneth L. Gentry was born in Springfield, IL, on July 24, 1977. He received a B.S.
degree in biomedical engineering from Marquette University, Milwaukee, WI, in 1999.
Currently, he is a biomedical engineering doctoral candidate at Duke University,
Durham, NC. He currently researches the design and fabrication of intracardiac transducers for both real-time volumetric imaging and
ultrasound ablation.
Stephen W. Smith (M’91) was born in Covington, KY, on July 27, 1947. He received the
B.A. degree in physics in 1967 from Thomas
More College, Ft. Mitchell, KY, the M.S.
degree in physics in 1969 from Iowa State
University, Ames, and the Ph.D. degree in
biomedical engineering in 1975 from Duke
University, Durham, NC.
In 1969, he became a Commissioned Officer in the U.S. Public Health Service, assigned to the Food and Drug Administration,
Center for Devices and Radiological Health,
Rockville, MD, where he worked until 1990 in the study of medical
imaging, particularly diagnostic ultrasound and in the development
of performance standards for such equipment. In 1978, he became
an adjunct associate professor of radiology at Duke University Medical Center. In 1990, he became associate professor of biomedical
engineering and radiology, and Director of Undergraduate Studies in
Biomedical Engineering at Duke University.
Dr. Smith has served on the Education Committee of the American Institute of Ultrasound in Medicine, the Executive Board of the
American Registry of Diagnostic Medical Sonographers, and the Editorial Board of Ultrasonic Imaging. He was co-recipient of the American Institute of Ultrasound in Medicine Matzuk Award in 1988 and
1990 and co-recipient of the IEEE-UFFC Outstanding Paper Award
in 1983 and 1994. He holds seven patents in medical ultrasound and
has authored 100+ publications in the field.