A_Thin_Transducer_With_Integrated_Acoustic_Metamaterial_for_Cardiac_CT_Imaging_and_Gating.pdf

1064 IEEE TRANSACTIONS ON ULTRASONICS, FERROELECTRICS, AND FREQUENCY CONTROL, VOL. 69, NO. 3, MARCH 2022
A Thin Transducer With Integrated Acoustic
Metamaterial for Cardiac CT
Imaging and Gating
Stephan Strassle Rojas , Graduate Student Member, IEEE, Srini Tridandapani , Senior Member, IEEE,
and Brooks D. Lindsey , Member, IEEE
Abstract— Coronary artery disease (CAD) is a leading
cause of death globally. Computed tomography coronary
angiography (CTCA) is a noninvasive imaging procedure for
diagnosis of CAD. However, CTCA requires cardiac gating
to ensure that diagnostic-quality images are acquired in all
patients. Gating reliability could be improved by utilizing
ultrasound (US) to provide a direct measurement of cardiac
motion; however, commercially available US transducers
are not computed tomography (CT) compatible. To address
this challenge, a CT-compatible 2.5-MHz cardiac phased
array transducer is developedvia modeling, and then, an ini-
tial prototype is fabricated and evaluated for acoustic and
radiographic performance. This 92-element piezoelectric
array transducer is designed with a thin acoustic backing
(6.5 mm) to reduce the volume of the radiopaque acoustic
backing that typically causes arrays to be incompatible with
CT imaging. This thin acoustic backing contains two rows of
air-filled, triangular prism-shaped voids that operate as an
acoustic diode. The developed transducer has a bandwidth
of 50% and a single-element SNR of 9.9 dB compared to
46% and 14.7 dB for a reference array without an acoustic
diode. In addition, the acoustic diode reduces the time-
averaged reflected acoustic intensity from the back wall
of the acoustic backing by 69% compared to an acoustic
backing of the same composition and thickness without the
acoustic diode. The feasibility of real-time echocardiogra-
phy using this array is demonstrated in vivo, including the
ability to image the position of the interventricular septum,
which has been demonstrated to effectively predict cardiac
motion for prospective, low radiation CTCA gating.
Index Terms— Acoustic diode, acoustic metamaterial,
computed tomography (CT), computed tomography coro-
nary angiography (CTCA), CT-compatible transducer.
Manuscript received December 2, 2021; accepted December 27, 2021.
Date of publication December 31, 2021; date of current version March 3,
2022. This work was supported in part by the Department of Biomed-
ical Engineering and the College of Engineering, Georgia Institute of
Technology; and in part by the National Science Foundation under Grant
ECCS-2025462. (Corresponding author: Brooks D. Lindsey.)
This work involved human subjects or animals in its research. Approval
of all ethical and experimental procedures and protocols was granted
by the local Institutional Review Board (IRB) at Georgia Institute of
Technology.
Stephan Strassle Rojas and Brooks D. Lindsey are with the
Georgia Institute of Technology, Atlanta, GA 30332 USA (e-mail:
brooks.lindsey.
@.
bme.gatech.edu).
Srini Tridandapani is with the Department of Radiology, The University
of Alabama at Birmingham, Birmingham, AL 35249 USA.
Digital Object Identifier 10.1109/TUFFC.2021.3140034
I. INTRODUCTION
CORONARY artery disease (CAD) accounted for 18%
of deaths in USA and as high as 46% of deaths in
some countries in 2019 [1]. The gold standard for diagnosing
CAD is catheter coronary angiography (CCA); however, CCA
is an invasive procedure and requires the use of a cardiac
catheterization laboratory and highly specialized staff, which
are not accessible in all parts of the world. Computed tomogra-
phy coronary angiography (CTCA) is a noninvasive, low-cost
alternative for imaging coronary arteries using X-ray computed
tomography (CT) imaging. CT imaging systems are much
more widely available relative to cardiac catheterization labs.
For example, in Lithuania, a country where CAD accounted
for 38% of deaths in 2019 [1], there are 20.2 CT systems
per million inhabitants [2]. In USA, 96% of emergency rooms
have access to CT, whereas only 36% of acute care hospitals
have catheterization labs [3], [4]. In developing nations, the
gap between the availability of CTCA and CCA is particu-
larly large. For example, in Senegal, there is only a single
catheterization laboratory for a population of 14.1 million
compared to five CT systems available in the public sector
alone [2], [5], [6].
In order to provide noninvasive diagnostic imaging via
CTCA, all CT systems must be able to acquire diagnostic-
quality scans, including older systems, which may be the
only systems available in many locations. Diagnostic-quality
CTCA images are acquired during the quiescent phase of the
cardiac cycle, i.e., when cardiac motion is at a minimum. Gat-
ing can be performed either retrospectively or prospectively.
In retrospective gating, CT images are acquired continuously,
with imaging data affected by a cardiac motion excluded
in postprocessing. In prospective gating, electrocardiography
(ECG) is used to trigger the acquisition of CT data, thus lim-
iting the patient’s exposure to radiation by only activating the
X-ray source during the quiescent phase of the cardiac cycle.
However, the viability of prospectively gated CTCA depends
on how effectively ECG can predict cardiac quiescence. ECG
gating is particularly challenged in patients with elevated heart
rates, high heart rate variability (HRV), or cardiac arrhythmias,
which can result in nondiagnostic CTCA images [7], [8]. Cur-
rently, only two-thirds of CTCA images are considered “good”
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STRASSLE ROJAS et al.: THIN TRANSDUCER WITH INTEGRATED ACOUSTIC METAMATERIAL FOR CARDIAC CT IMAGING 1065
quality, with a diagnostic accuracy of 80% [9]. A recent study
of 208 patients showed that the predictive value of CTCA
depends heavily on scan quality [9]. Newer CT systems with
multiple detector rows and decreased scan acquisition times
can reduce the effects of cardiac motion; however, diagnostic
performance is still challenged in some patients even with
these systems [10].
While ECG is a surrogate for cardiac motion based on the
assumption that electrical activity reflects mechanical motion,
direct sensing of cardiac mechanics can improve CTCA gating.
Cardiac monitoring with ultrasound (US) has been demon-
strated to accurately detect cardiac quiescence [11], [12],
with CTCA images from echocardiography-selected phases
having comparable quality to retrospective gating, but with
the advantage of significantly reducing radiation dose to the
patient [11], [12]. However, most currently available piezo-
electric US transducers are incompatible with CT systems due
to the presence of dense, radiopaque materials. Alternatively,
other groups described incompatibility of transducers for MRI
and have developed MR-compatible transducers [13]–[16].
Thus, placing a US transducer on the patient’s chest during
CT acquisition results in prominent streak artifacts in the
CT images preventing consistently reliable visualization of
coronary arteries [17], [18]. The materials in US transducers
that are incompatible with CT imaging include metals used for
interconnects or to shield the probe from electromagnetic inter-
ference, high-density piezoelectric materials, and high-density
acoustic backings, which can represent a large fraction of the
total volume of the transducer. While highly X-ray attenuating
metal components such as cables or interconnects can be
positioned outside the field of view of the CT system to mini-
mize the effects on resulting images, materials in the acoustic
stack cannot be relocated outside the field of view. We have
previously demonstrated that the acoustic backing typically
found in commercial US probes was a major source of the
streak artifacts when placed within the CT imaging field of
view [18]. These acoustic backings are often composed of
epoxies loaded with high-density powders to provide high
acoustic impedance and high attenuation coefficients, ensuring
the acoustic waves from the piezoelectric material propagate
into the acoustic backing, where they are then attenuated [18].
We also demonstrated a 20-fold improvement in CT compat-
ibility by designing a low-frequency (2.5 MHz) transducer
with an acoustic backing consisting of an epoxy loaded with
aluminum oxide (Al2O3) particles while maintaining sufficient
penetration depth for echocardiography [18].
While our previous study demonstrated reduced CT artifacts
by modifying the acoustic backing composition, even these
improved transducers have radiopacity similar to that of bone
and still result in some artifacts in CT scans [18]. In order to
further reduce artifacts and shadowing in CT images due to the
transducer, reducing the total volume of the acoustic backing is
required. However, the acoustic backing serves a critical pur-
pose for transducer performance, attenuating acoustic energy
that enters the backing before it can be reflected back into
the piezoelectric material. In order to attenuate reflections
at the back surface of the piezoelectric material and ensure
broad bandwidth (BW), thick acoustic backings (>10λ) with
high acoustic impedances (≥5.5 MRayl) and high attenuation
coefficients are typically used [19]. Previous work demon-
strated that an epoxy loaded with 25% v/v aluminum oxide
provides the attenuation of 2 dB/mm at 2.5 MHz and a cardiac
array transducer using this backing does not have noticeable
reflections from the back wall of the acoustic backing for
a 20-mm-thick acoustic backing [18]. While decreasing the
thickness of the acoustic backing would result in improved
CT compatibility, simply decreasing the thickness alone would
result in artifacts in B-mode images due to high-amplitude
reflections from the back surface of the acoustic backing,
which would not be adequately attenuated in a thin acoustic
backing.
In order to provide sufficient acoustic attenuation in a small
space, several approaches have been demonstrated. A phase
cancellation backing structure that splits acoustic energy into
two separate acoustic backings composed of different materials
has been developed [20], [21]. The differences in acoustic
properties between the two materials result in different inter-
actions at the back wall of the acoustic backing, with one
material producing a phase inverted reflection, while the other
material produces a noninverted reflection. This phase inver-
sion results in destructive interference when the two waves
in the different materials meet at the front of the backing.
While this approach enables thinner backing, the phase cancel-
lation backing structure uses a complex multimaterial backing
designed for a single center frequency. Alternatively, other
researchers have demonstrated modification of the back wall of
the acoustic backing to promote oblique back wall reflections,
resulting in longer wave propagation paths and increased
attenuation of the acoustic energy [22]. This approach is not
limited by the acoustic properties of the backing, and however,
it still relies heavily on having a sufficiently thick acoustic
backing to attenuate the oblique reflections before reaching the
interface between the piezoelectric material and the acoustic
backing. Furthermore, while thin US transducers have been
demonstrated, they are challenged to achieve the penetration
depth needed for cardiac imaging [23]–[25] and incorporate
non-CT compatible components such as electronic circuitry
and interconnects in the CT field of view [25].
Alternatively, acoustic echoes can be prevented from prop-
agating back into the piezoelectric material by designing
an acoustic backing containing structures to direct echoes
away from the piezoelectric material and confine acoustic
propagation within an attenuating material. In this way, thin-
ner backing with similar attenuation capabilities to its much
thicker counterpart can be achieved.
In this article, the development of a 2.5-MHz cardiac
phased array transducer with an acoustic metamaterial backing
structure for improved radiographic compatibility is described.
This thin transducer could be used to prospectively gate the
acquisition of CTCA images. First, a lossy acoustic back-
ing with an acoustic diode integrated into the backing was
designed using FEA modeling. Next, the designed device
was fabricated, resulting in a 92-element cardiac phased array
transducer. Finally, both the acoustic and radiographic perfor-
mances of this array were evaluated in phantom and in vivo
imaging. To our knowledge, this is the first development

Fig. 1. Illustrative diagram (not to scale) of the acoustic diode structure
used in the designed transducer is shown, consisting of two rows of
triangular voids in a lossy material (Al2O3 + epoxy). Diode parameters
shown are as follows: total backing thickness (HBacking), spacing between
PZT and diode (αtri), triangle height (htri), spacing between triangle rows
(βtri), and triangular void vertex angle (θ).
of a medical US transducer with an integrated acoustic
metamaterial.
II. METHODS
An acoustic diode is a metamaterial structure that was only
first experimentally demonstrated in the past two decades [26].
For an acoustic diode design to be effective in this application,
it will be necessary for the diode to have a high transmission
coefficient in the forward direction while maintaining a low
transmission coefficient in the reverse direction. While there
are many proposed acoustic diode designs, those designs
with a low transmission coefficient in the forward direction
are unsuitable for this application [27]–[29]. In addition, the
acoustic diode must operate over a sufficiently broad BW for
a US imaging transducer and must be able to be manufactured
within a thin transducer backing. Other groups have previously
demonstrated the feasibility of acoustic backings with complex
internal structures [30]. With these requirements in mind,
an acoustic diode design comprised of periodic, triangular
prism-shaped voids were pursued, following the previous
design of an acoustic diode at audio frequency (8.95 kHz),
demonstrating both broad BW and a high transmission coef-
ficient in the forward direction [31]–[33]. A similar acoustic
performance is desired at 2.5 MHz for a thin, CT-compatible,
cardiac phased array transducer.
A. Broadband Acoustic Diode: Theory of Operation
Previous work using finite difference simulations demon-
strated the use of closely packed, periodic equilateral triangular
structures to produce a highly efficient acoustic diode [31].
The geometry used in this design, which consisted of two
rows of repeating triangular voids, is shown in Fig. 1. In the
previous simulations with this design, the authors reported that
the efficiency of the diode was heavily affected by the size of
the repeating triangular structure relative to the wavelength
of the acoustic wave [31]. The operation of the acoustic
diode was not described by closed-form equations, but rather
was demonstrated through numerical simulations due to the
complexity of interference when the void geometry is similar
to the wavelength of the incoming wave [34].
Building on the theory developed in this previous work [31],
another group demonstrated the experimental feasibility of
a periodic triangular structure for operating as a broadband
diode at audible frequencies using eight rows of wooden
triangular prism pillars [33]. Implementing an acoustic backing
at ultrasonic frequencies introduces new sets of considera-
tions. The geometric parameters, material properties, and the
position of the acoustic diode within the acoustic backing
all determine the performance of the acoustic diode. Five
geometric parameters that determine the performance of the
acoustic diode are shown in Fig. 1: the spacing between
the PZT and the diode (αtri), triangle height (htri), spacing
between triangle rows (βtri), triangular void vertex angle (θ),
and spacing between the acoustic diode and the back wall of
the acoustic backing (HBacking).
In addition to the aforementioned parameters, the width of
the triangular voids is determined by the height of the triangle
and the vertex angle. The width of the triangle then determines
the void pitch (Ptri), which is twice the width of the triangles.
The effects of most parameters on the performance of the
acoustic diode were unknown at the start of the study, given
that only htri and βtri were varied in the previous publication
describing the diode structure in simulations [31]. In these
previous simulations, the BW of the acoustic diode depended
on the size of the voids relative to the design center frequency.
In addition, previously published simulations of the periodic
acoustic diode structure also showed that the spacing between
triangle rows (βtri) determined the efficiency of the diode (high
forward transmission and low reverse transmission) [31].
While the effects of αtri, θ, and HBacking are unknown
based on previous publications, the effects of each can be
hypothesized. For both αtri and HBacking, high values result
in increased thickness of the lossy acoustic backing, thus
increasing attenuation. Increasing the vertex angle θ is likely
to reduce the effectiveness of the acoustic diode structure.
For example, as θ is increased to 180◦
, then the transmission
coefficient in the forward direction approaches 0.
B. Simulations
To test the viability of adapting the highly efficient acoustic
diode design in [31] for use as an acoustic backing in a medical
US transducer at 2.5 MHz, a 2-D acoustic simulation model
was developed using COMSOL Multiphysics, Stockholm,
Sweden. This model simulates the behavior of a lossy acoustic
backing that incorporates the acoustic metamaterial (diode)
design. Two designs were modeled: a design with an acoustic
diode integrated within the lossy backing (30% Al2O3 + EPO-
TEK 301 epoxy) and a reference design containing only a
uniform, homogeneous lossy backing (30% Al2O3 + EPO-
TEK 301 epoxy) without the integrated acoustic diode. Models
of both designs used a Rayleigh damping model to simulate
the lossy properties of the backing material. The acoustic
properties of the loaded epoxy were characterized in our
previous work [18]. The acoustic diode design within the

TABLE I
DIODE SIMULATION PARAMETERS
backing consists of two rows of isosceles triangles, which were
left as voids with free boundary conditions to represent the air-
filled pockets found in the physical acoustic diode. The initial
dimensions for these voids were determined according to the
dimensionless simulation results in [31] to maximize the BW
of the acoustic diode at a center frequency of 2.5 MHz. These
base values were found by scaling values in [31] to the design
center frequency of 2.5 MHz and are shown in Table I. The
values of dimensions that are unique to this acoustic diode
design that is integrated into an acoustic backing (HBacking
and αtri) were chosen so that the diode would be located in
the middle of the backing material while keeping the backing
thickness less than 10 mm. The design performance was tested
by performing time-domain simulations of the model in which
a 0.75-cycle, Gaussian-windowed plane wave was excited at
the front of the backing model. The length of the model
duration in time was sufficient to allow multiple reflections
in all structures and ensure that the majority of the acoustic
energy had been dissipated (typically ∼55 μs).
To ensure the desired transducer performance, the prop-
erties of the triangular pillars were varied from the initial
values obtained in [31]. Multiple parameters of the diode
array structure were varied in these investigations, starting
with the placement of the two rows of triangular structures
within backing. The position of these rows is determined by
parameters αtri and HBacking, which are the spacing between
the PZT and the diode, and the spacing between the diode
and the back wall of the acoustic backing, respectively. Next,
the separation between rows of triangular structures was varied
in the depth direction (βtri). The effect of varying the vertex
angle of the air-filled triangular voids (θ) was also investigated.
Because a constant void width was used, varying θ also
resulted in variation in void height. In addition, simulations
in which void height (htri) was varied while maintaining a
constant vertex angle were also performed. Because varying
void height with a constant vertex angle results in variation of
void width, this case also resulted in the variation of void
pitch (Ptri). These geometric parameters, shown in Fig. 1,
were varied across the range of values described in Table I.
These ranges were selected by first determining a wavelength
to void height ratio that demonstrated high transmission in
the forward direction and low transmission in the backward
direction based on previous acoustic diode designs [31]. Next,
geometric parameters were scaled so that the design frequency
of the diode would fall between 0.75 and 6 MHz, which
is beyond the frequency range of operation of the designed
transducer. In addition, some parameters (HBacking, αtri, and θ)
are specific to the design of an acoustic backing and thus
were not studied in previous investigations of acoustic diode
operation [31]. These values were assigned a simulation range
of −50% to +250% of the base value determined by scaling.
An upper limit of +250% was chosen to limit the thickness
of the acoustic backing to ∼10 mm. A lower limit of −50%
was chosen to ensure that the designed acoustic backing could
be fabricated. In addition, only one parameter was varied in
each simulation, with all other parameters remaining fixed to
their base design values (Table I).
The performance of the acoustic backing in reducing reflec-
tions from the back surface of the acoustic backing was
assessed by measuring the time-averaged acoustic intensity
across the backing–piezoelectric interface using the following
equation:
Ir
Ii
(1)
where Ii is the acoustic intensity of the incident wave crossing
the backing–piezoelectric interface and Ir is the acoustic
intensity reflected by the backing structure crossing the same
interface. The angle brackets in (1) indicated that the time
average of the value was computed using an averaging window
equivalent to the total duration of the simulation.
Measuring Ir is important because high-amplitude
acoustic waves traveling from the acoustic backing into the
piezoelectric material would produce a strong artifact dis-
rupting the US image and potentially rendering it useless.
Moreover, high values of reflected time-averaged acoustic
intensity also result in a transducer having longer ringdown,
i.e., lower BW. Reflected time-averaged acoustic intensity
is also a good indicator of general backing performance
because reflected power at the acoustic backing–piezoelectric
interface results from acoustic energy that is neither trapped
by the diode structure nor attenuated by the lossy backing
material. Because the acoustic wave incident on the front of
the piezoelectric material is the sole source of acoustic energy
in the model, by normalizing Ir with respect to Ii , the
fraction of acoustic intensity attenuated by the backing can
be determined. By comparing Ir /Ii across acoustic diode
simulations with different parameter values (Table I) as well
as with the homogeneous reference acoustic backing of equiv-
alent thickness, it was possible to design an acoustic backing
capable of minimizing ringdown and in turn increasing BW
while also reducing total acoustic backing thickness to improve
CT compatibility.
C. Fabrication
Following modeling-based determination of parameters in
the acoustic diode design (Fig. 1 and Table I), a physical

Fig. 2. Multistep casting process was used to embed triangular prism-
shaped air-filled voids into a resin casting to generate an acoustic diode
within a thin, lossy acoustic backing for a phased array transducer.
phased array transducer was fabricated to test the feasibility
of the thin transducer design with an integrated acoustic diode
backing. The transducer consisted of 92 elements of PZT-5H
(HK1HD, TRS Technologies, State College, PA, USA) with
a height of 10 mm (elevation), a width of 0.18 mm, and
a thickness of 0.76 mm. The interelement separation was
0.280 mm. The PZT was bonded onto a custom polyimide
flexible interconnect using conductive epoxy (E-solder 3022,
vonRoll, Breitenbach, Switzerland). A 270-μm-thick matching
layer (∼λ/4) was bonded to the front of the PZT elements.
To maintain radiographic compatibility for CT imaging when
the transducer is placed on the patient’s chest, both the
matching and backing layer were composed of an epoxy
(EPO-TEK 301, Epoxy Technology, Billerica, MA, USA)
loaded with Al2O3 particles (15 μm, 50362-15, Electron
Microscopy Sciences, Hatfield, PA, USA) at a concentration
of 30% by volume, as in modeling [18]. For the reference
transducer, a 0.65-cm-thick lossy backing was attached to the
back of the interconnect, while the diode transducer contained
the described metamaterial backing.
The backing structure was fabricated utilizing a multistep
epoxy casting process to generate the triangular prism-shaped
voids required to create the acoustic diode within the acoustic
backing. The steps used to generate the acoustic diode backing
are shown in Fig. 2. Each casting step utilized a two-piece
polyacetal negative mold that was micromachined by a CNC
mill (HAAS CM-1). The molds had an opening at the top for
the pouring of uncured resin. The resin was poured into the
first set of molds and then allowed to cure for the first casting.
Once cured, the cast resin was removed by separating the two
halves of the mold, a step that was repeated after each casting.
The subsequent casting steps utilized mold negatives that could
accommodate the previous castings within the mold cavity.
This enabled the original casting to be built up to the final
size using subsequent castings, where each new casting added
a new material layer that contained layer-specific part features,
as shown in Fig. 2. This multicasting technique is similar
to overmolding, which is an injection molding process [35].
During the first two casting steps (Fig. 2, Steps 1 and 3),
one of the mold halves imprinted the triangular prism void
structure onto the backing, leaving triangular channels once
the molds were released. These channels were then filled
with a water-soluble wax (Sol-U-Carv, Freeman, Avon, OH,
USA), which prevented resin from filling these channels in
the subsequent casting steps and thus served as a lost-wax
core (Fig. 2, Steps 2 and 4). Once all casting steps had been
completed, the wax was dissolved in a 70 ◦
C water bath in
an ultrasonic cleaner for 90 min, leaving triangular prism-
shaped voids running the length of the acoustic backing in the
elevation direction. This process ensured that the triangular
prisms maintained a high degree of dimensional tolerance.
The homogeneous backing used in the reference array and
the matching layers in both arrays were also produced using
two-piece, polyacetal micromachined molds. However, only a
single casting step was required to produce each of these parts.
The transducer with the acoustic diode backing was con-
structed by bonding a 0.76 mm × 10 mm × 27 mm piece
of PZT with 100 nm of gold-sputtered on the top and bottom
surfaces to a flexible polyimide interconnect. The individual
elements were then separated using a dicing saw (ADT 7100,
Advanced Dicing Technologies, Zhengzhou, China) with a
100-μm kerf dicing blade (4B776-3AB1-040-BL0, Advanced
Dicing Technologies, Zhengzhou, China). The top surfaces of
the PZT elements were then shorted together with a bead of
conductive epoxy placed along the outer edge of the elements
in the elevation direction to form a uniform ground electrode
for all elements without obstructing the front surface of the
elements. On the opposite (back) side, the custom flexible
interconnect created an independent signal electrode at the
bottom of each PZT element. At this point, the preformed
matching layer was bonded onto the top of the electrode,
whereas the preformed backing layer described above was
bonded to the back of the interconnect. Both the matching
layer and the acoustic backing were bonded using a thin
layer of EPO-TEK 301 epoxy loaded with 30% Al2O3 to
create bonding layers having the same acoustic impedance
as the passive acoustic layers. The transducer was finished
by coating with 10 μm of Parylene to prevent electrical
shorts and protect the device from the external environ-
ment. The thickness of the completed backing attached to
each transducer was then measured using a digital caliper
(500-151-30, Mitutoyo, Kawasaki, Japan).
The completed array was connected to a research US system
(Verasonics Vantage 256, Kirkland, WA, USA) via a custom
cable between the flexible interconnect and a zero insertion
force connector (DL5-260PW6A, ITT Cannon, Irvine, CA,
USA). This cable was attached to an intermediary custom
printed circuit board, which was then connected to the flexible
interconnect on the transducer. The transducer was connected
to the system by approximately 5 ft (1 m) of cabling, thus
requiring only the transducer to be positioned within the CT
system. The distal end of the cable, intermediary circuit board,
and transducer were then housed in a 3-D printed plastic
enclosure to protect the assembly as well as improve the

Fig. 3. (a) Completed phased array US transducer with cable assembly
and 3-D printed enclosure is shown. (b) Front (end) view of the array
transducer acoustic stack shows triangular voids in the acoustic backing.
(c) Magnified view of the void structure in the acoustic backing.
ergonomics of using the device with a patient lying in the
CT scanner while maintaining CT compatibility. The fully
assembled transducer is shown in Fig. 3.
D. Transducer Characterization
Transducer acoustic performance was evaluated through
single-element pulse-echo testing of 15 elements in both
the transducer containing the acoustic diode in its acoustic
backing and the reference transducer without the acoustic
diode. Testing was conducted in a water tank, with the tip
of a 254-μm-diameter wire at a distance of 10 mm from the
transducer element serving as a point target for the element
being tested. A pulser–receiver (Panametrics 5073PR, Olym-
pus, Waltham, MA, USA) was used to excite an individual
element in the transducer (−190 V impulse and 10-ns fall
time). The received radio frequency (RF) echoes were digi-
tized at 100 MHz using a 14-bit acquisition board (Signatec
PDA14, Corona, CA, USA). The acquired waveforms were
processed using MATLAB (Mathworks, Natick, MA, USA).
The SNR was calculated by measuring the root-mean-square
(rms) amplitude of the noise in front of each pulse-echo
waveform. This was divided by the rms amplitude of the pulse-
echo waveform for each element.
Radiographic properties of the transducer were evaluated
using a two-step process. First, the transducer was scanned
with a microCT system (μCT50, Scanco Medical, Brüttisellen,
Switzerland). For this scan, only the transducer array was
imaged, with no housing or cabling to evaluate the radi-
ographic performance of the acoustic stack. The transducer
was then positioned on a CT chest phantom and scanned with
a clinical CT system, this time with both the 3-D printed
ergonomic enclosure and the cable to evaluate the compati-
bility of the completed device for clinical CT gating.
E. US Imaging
The US imaging performance of both the fabricated diode
and reference transducers was evaluated by acquiring B-mode
US images using a tissue-mimicking phantom (ATS Labs
Model 539, Bridgeport, CT, USA, α = 0.5 dB cm−1
MHz−1
)
having a series of 0.25-mm wire targets embedded at varying
depths ranging from 5 to 80 mm. Spatial resolution was
evaluated by measuring the cross section of a wire target in
the tissue-mimicking phantom at a depth of 4.8 cm using
RF data prior to log compression for image display. The
transducer with the metamaterial acoustic backing and the
Fig. 4. Simulated pressure fields resulting from a short 2.5-MHz plane
wave incident on the diode structure. (a) At T = 1.1 μs, a plane wave
propagates through the backing uninterrupted. (b) At T = 2.2 μs into
the simulation, the plane wave begins to interact with the first row of
triangular voids, resulting in most acoustic energy propagating past the
first row of voids, however, some energy is reflected back toward the
source of the wave. (c) At T = 2.9 μs into the simulation, the wave starts
to interact with the second row of voids. Some of the acoustic energy
keeps traveling through the diode, while some is trapped between the
two rows of triangular voids, and the remaining energy leaks through the
first row and propagates toward the source of the wave. (d) At T = 3.6 μs
into the simulation, the energy remaining in the original plane wave starts
to interact with the back surface of the acoustic backing and is reflected
back toward its source. Most of this reflected energy will be trapped by
the acoustic diode, while a small fraction will leak through and arrive back
at the front (upper) surface of the acoustic backing.
reference transducer with the conventional, isotropic acoustic
backing were both used to image the phantom.
Finally, to demonstrate the proof of concept of the fabricated
transducers, in vivo cardiac imaging was performed in an
apical four-chamber view in one healthy volunteer via IRB-
approved protocol using the transducer with the acoustic diode
backing. To mimic the intended use of these transducers for
cardiac gating in a CT system, imaging data were acquired
over multiple cardiac cycles, allowing both B- and M-mode
data to be extracted from acquired datasets. B-mode data
were postprocessed by applying a median filter with a ker-
nel size of 2.5 mm × 1.5 mm (lateral × axial) to reduce
the effect of noise found within the image. The M-mode
data were upsampled from 29 samples/s × 3.3 samples/mm
to 29 samples/s × 26.6 samples/mm (time × axial), and then,
the data were filtered with a median filter having a kernel size
of 34 ms × 0.1 mm (time × axial).
III. RESULTS
A. Simulation
The geometric parameters in the acoustic diode structure
were varied individually across the ranges of values in Table I.

Fig. 5. (a) Pressure field within the backing for the simulated case
when vertex angle θ = 70◦
is shown. (b) Magnified view of simulated
pressure fields within the acoustic backing reveals a spatially varying
pressure pattern at the PZT–backing interface with locations of zero
acoustic pressure. For acoustic diode parameters that differ from those
shown here, these locations of zero acoustic pressure would occur at
different depths relative to the PZT–acoustic backing interface.
In Fig. 4, the behavior of the wavefront at four different
stages of interaction with the diode structure is shown: the
uninterrupted wavefront [Fig. 4(a)], the wavefront’s first inter-
action with the initial row of voids [Fig. 4(b)], the wavefront’s
first interaction with the second row of voids [Fig. 4(c)],
and finally the wavefront’s first interaction with the back
wall of the acoustic backing [Fig. 4(d)]. In Fig. 5, the gen-
eration of spatial locations of high and low pressure at the
PZT–backing interface can be observed, including locations
where acoustic pressure is close to zero. These peaks and their
locations depend on void dimensions and their offset from the
PZT–backing interface. For the case of a 70◦
vertex angle,
as shown in Fig. 5, these peaks occur at the PZT–acoustic
backing interface itself. This is likely the result of the effect of
the vertex angle on the interference, as varying the vertex angle
changes the locations of peak positive and negative acoustic
pressure. Fig. 6(a)–(e) shows the effect of varying several
geometric parameters in the acoustic diode backing design on
time-averaged acoustic intensity. Alternatively, Fig. 6(f) shows
the effect of varying the total thickness of the homogeneous
backing on the reflected acoustic intensity for the reference
array with a conventional, lossy backing only. The exponential
decay observed in Fig. 6(f), which is in log scale, implies that
beyond a certain thickness, further increasing the thickness
of the acoustic backing does not provide additional benefits
because any acoustic reflections from the back wall will be
below the noise floor and thus will not have a meaningful
effect on performance.
Simulation results (Fig. 6) indicate that the key parameters
that determined the performance of the acoustic diode backing
were the vertex angle of the triangular prisms [θ, Fig. 6(b)]
and the offset between the rows of triangular air-filled voids
and the piezoelectric material [αtri, Fig. 6(d)]. While the spac-
ing between the rows of triangular voids (βtri) and the space
behind the diode (HBacking) has some influence on the perfor-
mance of the diode [Fig. 6(a) and (c)], this effect was small,
0.5 dB for the values tested. This observation suggests that
the acoustic diode was effective in confining acoustic energy
between the two rows of voids. In addition, as shown in
Fig. 6, HBacking [Fig. 6(a)] and βtr [Fig. 6(c)] primarily affect
the attenuation performance when their values are small. This
trend means that the values for spacing between rows of
voids (βtri) and the space behind the diode (HBacking) are the
primary parameters that can be minimized to further reduce
the thickness of the backing layer.
A simulation analysis resulted in the selection of a final
diode design consisting of triangular voids with θ = 30◦
vertex angles, triangle base width of 696 μm, and height of
htri = 1300 μm. Based on simulations, the selected design
has two rows of triangular prism voids with a spacing of
βtri = 500 μm, with the front row of voids positioned αtri =
3 mm behind the backing–PZT interface, whereas the back
surface of the acoustic backing is offset HBacking = 400 μm
from the back of the acoustic diode. In total, the designed
acoustic backing thickness is 6.5 mm. Given these results,
an acoustic backing with a thickness of 6.5 mm was used
for the fabricated transducers having both the acoustic diode
and homogeneous acoustic backings. With these parameters,
according to simulation results, a homogeneous backing of
the same thickness, 6.5 mm, would result in a reflected
time-averaged acoustic intensity of −20.1 dB, compared to a
reflected acoustic intensity of −25.2 dB for the acoustic diode
backing. According to simulations, a homogeneous backing
without an acoustic diode would need to be at least 8.5 mm
thick to perform similar to the diode containing backing.
B. Transducer Characterization
Based on the single-element acoustic characterization of the
two fabricated array transducers (the array with an acoustic
diode backing and the array with the reference backing), the
transducer with the integrated acoustic metamaterial backing
had a single-element SNR of 9.9 ± 2.1 dB for a point target
at a depth of 10 mm and a −6-dB BW of 50% ± 7.4%.
In comparison, the reference array with the homogeneous
(nonacoustic diode) acoustic backing had an SNR of 14.7 ±
1.1 dB and −6-dB BW of 46.7 ± 7.2%. The pulse-echo
waveform and the power versus frequency of a typical ele-
ment for both the transducer with the diode as well as the
reference transducer without the acoustic diode are shown in
Fig. 7. MicroCT testing revealed that both transducers had
radiopacities of ∼1200 HU, significantly lower than that of a
commercial tungsten-filled backing, which has a radiopacity
of 15 200 HU [18]. Clinical CT testing demonstrated the
transducer with the acoustic diode backing produced fewer
CT artifacts than its commercial counterpart (Fig. 8).
C. US Imaging
In imaging the commercial tissue-mimicking US phantom,
the transducer containing the acoustic diode backing exhibited
a spatial resolution of 1.59 mm × 0.81 mm (lateral × axial)
at a depth of 5 cm. In comparison, the reference array with
the homogeneous acoustic backing had a spatial resolution
of 1.73 mm × 0.83 mm (lateral × axial). Images acquired
using the two transducers as well as the lateral and axial cross
sections of the wire target used to calculate spatial resolution
are shown in Fig. 9.

Fig. 6. Normalized time-averaged acoustic intensity (Ir/Ii) measured in backings containing an acoustic diode in acoustic backing simulations as
a function of various geometric parameters. (a) Ir/Ii decreases slightly with increasing spacing between the back of the diode and the back wall of
the acoustic backing. (b) Ir/Ii reaches a minimum for small vertex angles, with a second local minimum occurring at ∼70◦. (c) Ir/Ii decreases
slightly with increasing spacing between the two rows of voids that comprise the diode. (d) Ir/Ii decreases with increasing spacing between the
front of the diode and the front wall of the backing. (e) Ir/Ii decreases as the height of the triangular void increases. However, the behavior is
nonlinear, with a local minimum at 500 μm. (f) Ir/Ii for the homogeneous backing decreases with increasing thickness for the reference transducer
without the acoustic diode. Because the homogeneous backing does not contain any structures within the backing, all the reflected acoustic energy
observed at the backing–PZT interface is the result of reflection from the back surface of the acoustic backing.
Fig. 7. Single-element pulse-echo waveform (shown in black) and power
versus frequency (shown in red) for the array transducer with (a) acoustic
diode backing and (b) reference transducer.
Finally, in vivo images acquired in a healthy adult volunteer
in an apical four-chamber view using the transducer with the
acoustic diode are shown in Fig. 10. Cardiac M-mode data
are also shown for a duration of 4 s to show the ability to
track cardiac motion over multiple cycles, which is useful
for identifying cardiac quiescence (Fig. 11). For the M-mode
data shown, a spatial location that crosses the aortic valve was
selected to illustrate the sensitivity to motion of a thin, highly
mobile structure.
IV. DISCUSSION
A. Simulation
A CT-compatible phased array transducer for cardiac imag-
ing and CT gating with an integrated acoustic metamaterial
backing was developed based on simulations, and its feasibility
was evaluated in phantom and in vivo imaging studies. Sim-
ulations demonstrated that the acoustic diode backing trapped
acoustic energy between the two rows of triangular prisms.
Eventually, the trapped acoustic energy was attenuated by the
lossy material comprising the backing.
The behavior observed when varying the spacing between
the rows of triangular prism-shaped voids (βtri) indicates that
there is little effect on attenuation performance when the
Fig. 8. Clinical CT image of (a) Philips C9-2 commercial US transducer
and (b) developed transducer with the acoustic diode backing. Both
images were obtained using a 100-kV source Philips Brilliance 64 Slice
CT scanner with an RSD RS-111 Anthropomorphic Thorax Phantom.
spacing is ≥3 mm. However, decreasing spacing below 3 mm
results in increased reflected time-averaged acoustic inten-
sity, showing a clear trend, although the increase is gradual.
In examining simulation results, this behavior is primarily
due to increased interaction between acoustic energy and the
vertices of the triangular voids that act like point scatterers.
As the spacing between rows of triangular voids increases,
less acoustic energy reflected in the forward direction by the
first row of voids interacts with the vertices of the second
row of voids. This effect results in more acoustic energy

Fig. 9. B-mode images of a tissue-mimicking US phantom (ATS
Labs 539) with wire targets acquired using (a) diode-backed transducer
and (b) reference homogeneous-backed transducer. Both images are
displayed with a dynamic range of 65 dB. (c) Lateral and (d) axial cross
sections of a wire target at a depth of 5 cm (indicated by yellow circle
in A) are shown for both the acoustic diode and the reference transducer.
Fig. 10. (a) In vivo US image acquired in an apical four-chamber view
using the developed array transducer with the acoustic diode backing.
The image is displayed with a dynamic range of 45 dB. (b) Corresponding
anatomy of the apical four-chamber view seen in the image in (a) is shown
with all chambers labeled.
interacting with the side of the second row of voids, thus
causing more energy to propagate toward the back of the
backing as intended.
For this reason, when the row spacing is increased beyond
3 mm, the effect is minimal, as the primary source for the
scattered acoustic energy in the second row of void vertices
at that point is from the original, forward-traveling plane
wave. Any benefit of increasing βtri beyond 3 mm is due
to increased attenuation resulting from the increased backing
thickness (although at the cost of reduced CT compatibility).
In addition, this point-like scattering effect of the void vertex
can also explain why the increase in reflected energy is small
even with minimal void spacing because vertices acting as
point scatters still result in spatial spreading of acoustic energy.
This spreading gives the backing material an increased space
over which to attenuate the acoustic energy.
In addition, the effect of varying the spacing between
the diode structure and the acoustic backing–piezoelectric
material interface [αtri, Fig. 6(d)] shows a clear tread of
exponential decay, with increased spacing resulting in a
reduction in reflected time-averaged acoustic intensity. This
trend is similar to the trend observed for the reference
array with the homogeneous backing [Fig. 6(f)]. Increasing
αtri resulted in reduced normalized time-averaged acoustic
intensity (Ir /Ii ), because when spacing is increased, all
acoustic energy is forced to travel through a longer path length
within the lossy material, resulting in increased attenuation.
The purpose of this front layer is to attenuate the acoustic
energy that was reflected from the diode. Thus, it only
affects the performance of the backing, not the effectiveness
of the diode in trapping acoustic energy. For this reason,
maximizing αtri will always result in backing with increased
attenuation, and however, increased thickness decreases CT
compatibility.
Conversely, the void vertex angle (θ) directly influences the
diode’s effectiveness in trapping acoustic energy, as shown in
Fig. 6(b). In general, increasing the vertex angle to a larger,
more obtuse angle results in a less effective backing. This
occurs because as the vertex angle becomes larger, reflection
angles of the acoustic waves interacting with the void structure
become smaller, resulting in more energy being reflected
back toward the piezoelectric layer. At a sufficiently large
vertex angle, the triangular voids behave more like the back
wall of the backing, resulting in only normal incidence wave
reflections. However, in Fig. 6(b), it can also be observed that
this trend is not linear, with a local maximum at 50◦
and a
local minimum at 70◦
. In the simulation for the case of 70◦
,
acoustic pressure at some locations along the PZT-acoustic
backing material interface was approximately zero, resulting
in approximately zero reflected acoustic intensity at those
locations (Fig. 5). At angles other than 70◦
, this effect was
minimal. This behavior results from the periodicity of the void
structure, which leads to portions of the pressure field experi-
encing total destructive interference in part due to the Bragg
diffraction [34]. Only some θ values experience this effect
because the vertex angle determines the spatial interference
pattern, which affects the interference pattern of the diode.
For the majority of angles, most of the destructive interference
occurred outside the PZT-acoustic backing material interface,
so this behavior had little benefit for those values.
A similar nonlinear behavior is observed in Fig. 6(e), which
shows the effect of varying the triangular void height (htri).
In this case, a local minimum exists for a height of 500 μm
with a local maximum at htri = 750 μm. This behavior
results from the fact that the void height is related to three
parameters in the diode structure: 1) triangle void height;

Fig. 11. (a) M-mode data acquired using the transducer with the acoustic diode backing over a duration of 4 s in the location indicated by the green
line in (b) apical five-chamber view. The corresponding B-mode image for the two time points indicated by the yellow dashed lines in (a) are shown
for (c) T1 = 0.9 s and (d) T2 = 2.2 s. All images are displayed with a dynamic range of 45 dB.
2) void width; and 3) void pitch (separation), which varies
with void width. Because the periodicity of the voids varies
with void height, the effect of Bragg diffraction on the pressure
field also varies. As with vertex angle, certain void heights
result in destructive interference at the PZT–acoustic backing
material interface such that a local minimum occurs at this
interface. However, the primary effect of varying void height is
the overall variation in the thickness of backing. A larger void
height value results in a thicker backing, which forces acoustic
energy to travel further within the lossy backing, leading to
increased attenuation. For this reason, the tallest void height
tested results in the lowest time-averaged acoustic intensity
[Fig. 6(e)].
Based on the observed effects of the different geometric
parameters on acoustic backing performance, the geometric
parameters used in the developed array transducer balanced
acoustic attenuation with backing thickness. According to
simulations, a 6.5-mm-thick acoustic backing with the acoustic
diode structure produces a backing with a 69% reduction
in reflected time-averaged acoustic intensity compared to
a homogeneous backing of the same thickness. The fabri-
cated transducers confirm a reduction in reflected amplitude.
In Fig. 7, a reflection from the back surface of the acoustic
backing can be seen at the expected point in time in the pulse-
echo waveform. In contrast, in the array with the acoustic
diode Fig. 7(a), it is much more difficult to identify reflected
energy because it is more distributed in time, though the peak
amplitude of the reflected wave is 32% lower than that of
reference design.
Furthermore, the simulations revealed that a homogeneous
backing would need to be 30% thicker to match the perfor-
mance of the diode-containing backing. Due to the simulation
approach in which geometric parameters were varied individu-
ally, it is possible that the design selected for fabrication does
not represent the optimum or thinnest design.
While a broader simulation strategy may improve results,
this approach is also computationally expensive. For example,
varying four parameters across ten values each would result
in 104
individual time-domain simulations.
B. Experimental Results
A benefit of acoustic diode integration is that it can enhance
the acoustic backing’s intrinsic attenuation performance. In the
case of the results presented in this article, the acoustic diode
improved the performance of a backing composed of an
epoxy loaded with Al2O3. However, diode integration could be
performed in other lossy materials, provided that the difference
in acoustic impedance between the bulk backing material and
the voids is sufficiently high. If a backing material can be
effectively molded, then an acoustic diode structure similar
to that used in this article can be integrated into the backing
to enhance its performance. This may provide greater design
flexibility relative to other approaches for reducing backing
thickness [20], [21]. In specific applications such as the one
presented in this article (CT compatibility), there are strict
limitations on the types of materials that can be used, and thus,
it is challenging to produce a high-performing transducer with
additional material constraints.
In addition, the array with the acoustic diode exhibited
higher BW but decreased SNR relative to the reference array
without the acoustic diode. This slight increase in BW and
decrease in SNR is likely due to improved damping in the
acoustic diode. This behavior is consistent with the simulation
results, and however, modeling indicated a greater improve-
ment in BW due to the acoustic diode. This deviation from the
modeling is likely due to the model’s limitations; specifically,
the full acoustic stack was not modeled and the model is only
2-D, not 3-D. The mismatch in acoustic impedance between
the acoustic backing and the rest of the transducer may result
in additional reflections at this boundary that reduce BW.
In addition, the 2-D nature of the simulation model does not
fully capture the physical reality of the 3-D geometry, such as
out-of-plane or sidewall reflections that would increase pulse-
length. Finally, the fabricated transducer has slight deviations
in dimensions relative to the model geometry. Specifically,
the walls of the triangular voids had a small degree of
convexity, which was introduced during the mold production
process. This results in varying reflection angle, altering delay
times and increasing constructive interference and pulselength.

The void convexity also reduced the sharpness of the vertex
of the void, resulting in more acoustic energy being scattered
by the void vertex.
It may be possible to increase SNR by modifying the
parameters or material composition of the acoustic diode
backing. However, the primary goal in this development of
the initial transducer was to demonstrate the feasibility of
CT compatibility and acoustic performance (i.e., minimize
artifacts in both US and CT images) rather than optimizing
BW or SNR.
In the future, comparing the performance of fabricated
acoustic diode transducers having various geometric parame-
ters (Table I) and reference transducers with acoustic backings
of different thicknesses would allow for a better understand-
ing of the effects on performance (SNR and BW) between
transducers with acoustic diodes and those solely utilizing
lossy materials to achieve the acoustic attenuation. It may also
be possible to reduce the thickness further or generalize this
concept to other transducers by replacing the lossy material in
which the diode is embedded with a material having increased
attenuation, although this would also affect the density and
acoustic impedance.
Testing with a clinical CT system (Fig. 8) revealed that
the transducer with the acoustic diode reduced CT artifacts
compared to its commercial counterpart. The reduction in
artifacts results from both using a backing material with more
favorable CT properties and reducing the total volume of the
acoustic backing. Most of the remaining artifacts produced by
the transducer with the acoustic diode backing (Fig. 8) arise
from the PZT, as these artifacts are not visible in slices of the
CT scan containing only the backing but not the piezoelectric
material.
C. In Vivo Imaging
During in vivo imaging with the developed array contain-
ing the acoustic diode, it was possible to identify all four
chambers and the intraventricular septum (IVS) in an apical
four-chamber view (Fig. 10). The ability to visualize the IVS
indicates that the imaging performance of this transducer is
sufficient for cardiac gating in CTCA, as previous studies
have indicated that US imaging of the IVS can be used to
effectively predict periods of cardiac quiescence [36]. When
combined with the favorable CT imaging results seen in Fig. 8,
this suggests that the developed transducer may be acceptable
for CTCA gating. The M-mode data shown in Fig. 11 provide
an example of the type of signal that might be used to gate
CTCA acquisition, with the M-mode data in Fig. 11 showing
the motion of the aortic valve over multiple cardiac cycles.
D. Future Directions
It may be possible to improve the performance of the
acoustic diode design used in this work via global parameter
optimization of the model. In addition, alternative acoustic
diode structures that are more complex but more efficient
such as those using near-zero refractive index metamaterial
prisms or a combination of metasurfaces and photonic crystals
may enable a thinner backing structure [37], [38]. Finally,
the development of algorithms utilizing real-time cardiac data
to predict cardiac quiescence in future cardiac cycles is
needed to test the developed transducer’s viability thoroughly.
In addition, accompanying hardware that can interface with
the CT-compatible transducer developed in this work and the
CT system is needed to predict cardiac quiescence based on
US data and provide a trigger to prospectively gate CTCA
acquisition accordingly. This signal processing hardware could
incorporate recent developments in low-cost front-end design
architecture to reduce its complexity and enable US gating to
become a low-cost upgrade to existing CT systems [39], [40].
The development of a CT-compatible US transducer could
have broader effects in the field of radiology other than
improved CT gating. Simultaneous US and CT imaging pro-
vides several additional unique opportunities. In dynamic CT
myocardial perfusion imaging, the hemodynamic effect of
coronary artery stenoses is quantified based on the acquisition
of multiple sequential images acquired during contrast agent
administration [41]. However, the images required to develop
a contrast wash-in curve are acquired over several heartbeats
(typically 10–15 cardiac cycles during a 30-s breath hold [42])
and result in a dose of 5–10 mSv. Depending on the protocol
used to acquire these scans, if CTCA is also performed,
the patient may be subjected to a similar additional dose of
5–10 mSv. In these patients, prospective US gating may result
in significantly lower radiation dynamic perfusion imaging
from several heartbeats compared with retrospective gating.
Alternatively, PET-CT imaging provides simultaneous imag-
ing of metabolism and anatomy. Despite many applications
in other fields, such as oncology and neurology, cardiac
metabolic PET imaging with 2-deoxy-2-[18
F]fluoro-D-glucose
(FDG) is relatively limited in its clinical utility. Recently, a
small animal system for combined PET, CT, and US imaging
has been developed with the goal of identifying cardiac
PET imaging applications while using US data to correct
for cardiac motion, improving PET contrast and resolution
[43], [44]. Such a system could allow simultaneous assessment
of cardiac anatomy (including coronary stenosis) via CT,
wall motion via US, and metabolism via PET. While we
have previously demonstrated the significant negative effect
of the transducer on CT images [18], the developers of
the combined PET-CT-US system reported that the effect
of the transducer on PET was only a “modest impact on
gamma-ray attenuation” [44]. While X-ray attenuation due
to the transducer would be higher, for small animal imag-
ing, the high-frequency transducer could be positioned outside
of the PET-CT field of view, which may not be possible for
human imaging. Thus, in addition to CTCA gating, providing
simultaneous echocardiography via CT-compatible US trans-
ducers could enable low radiation dose dynamic myocardial
perfusion imaging in a single heartbeat and multimodality
imaging of cardiac anatomy, mechanics, and metabolism, with
US providing motion correction in humans, as demonstrated
in small animals [44].
V. CONCLUSION
A thin, CT-compatible US array transducer was developed
for cardiac imaging and gating inside of a CT scanner.
This CT-compatible cardiac phased array utilized an acoustic
metamaterial in the acoustic backing for the first time to

reduce the total thickness of the acoustic backing and thus
improve the device’s CT compatibility. Acoustic simulations
demonstrated that the diode structure in the acoustic backing
of the transducer resulted in improved acoustic attenuation
with reduced backing thickness, with some dependence on
the selection of geometric parameters of the acoustic diode.
The transducer developed on the basis of these simulations
was a 2.5-MHz, 92-element array with an acoustic backing
that was 6.5 mm thick. According to simulations, this design
resulted in 69% less acoustic power being reflected into
the piezoelectric elements from the backing compared to a
homogeneous backing of the same thickness. The fabricated
transducer containing an acoustic diode backing was used to
perform in vivo imaging of a human heart in an apical four-
chamber view, allowing visualization of the four chambers and
the interventricular septum, indicating that this transducer may
be useful for cardiac gating. In addition, testing with a clinical
CT system showed a significant reduction in CT artifacts
compared to a conventional transducer. US-gated acquisition
of CTCA could provide diagnostic-quality CT images for the
evaluation of CAD in all patients in all locations, including
patients with elevated or variable heart rates, those suffering
from cardiac arrhythmias, and those in rural locations.
ACKNOWLEDGMENT
Some of the work was performed at the Georgia Tech
Institute for Electronics and Nanotechnology, a member of the
National Nanotechnology Coordinated Infrastructure (NNCI).
The content is solely the responsibility of the authors and does
not necessarily represent the official views of the National
Science Foundation.
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Stephan Strassle Rojas (Graduate Student
Member, IEEE) received the B.S. degree
in mechanical engineering and electrical
engineering from the University of Florida,
Gainesville, FL, USA, in 2018. He is currently
pursuing the Ph.D. degree in electrical and
computer engineering with the Georgia Institute
of Technology, Atlanta, GA, USA.
His current research interests include
multimodality imaging, acoustic metamaterials,
forward-viewing intervascular ultrasound
(FV-IVUS), and transducer development.
Mr. Strassle Rojas received the Best Student Paper Award at the 2021
IEEE International Ultrasonics Symposium.
Srini Tridandapani (Senior Member, IEEE)
received the B.E. degree in electrical engineering
from Anna University, Chennai, India, in 1988,
the M.S.E.E. and Ph.D. degrees in electrical
engineering from the University of Washington,
Seattle, WA, USA, in 1990 and 1994, respec-
tively, the M.D. degree (followed by residency
training in radiology) from the University of Michi-
gan, Ann Arbor, MI, USA, in 2001, the mas-
ter’s degree in clinical and translational research,
and the M.B.A. degree from Emory University,
Atlanta, GA, USA, in 2012 and 2015, respectively.
After postdoctoral training in computer science at the University of
California at Davis, Davis, CA, USA, he was an Assistant Professor of
electrical and computer engineering with Iowa State University, Ames,
IA, USA. A board-certified radiologist, he completed clinical fellowships
in cardiothoracic imaging and abdominal imaging with Emory University.
He is currently a Professor and the Vice Chair of imaging informatics at
the Department of Radiology, The University of Alabama Birmingham,
Birmingham, AL, USA.
Brooks D. Lindsey (Member, IEEE) received
the B.S. degree in electrical engineering from
the University of Illinois at Urbana–Champaign,
Champaign, IL, USA, in 2007, and the Ph.D.
degree in biomedical engineering from Duke
University, Durham, NC, USA, in 2012. He com-
pleted postdoctoral training at the Joint Depart-
ment of Biomedical Engineering, The University
of North Carolina at Chapel Hill, Chapel Hill,
NC, USA, and North Carolina State University,
Raleigh, NC, USA.
In 2017, he joined the Wallace H. Coulter Department of Biomedical
Engineering, Georgia Institute of Technology, Atlanta, GA, USA, and
Emory University, Atlanta, as an Assistant Professor, where he directs
the Ultrasonic Imaging and Instrumentation Laboratory. His research
interests include interventional imaging and development of ultrasound
transducers and systems.

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