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. Author manuscript; available in PMC: 2023 Jun 1.
Published in final edited form as: IEEE Trans Ultrason Ferroelectr Freq Control. 2022 May 26;69(6):1917–1925. doi: 10.1109/TUFFC.2022.3153929

Ultrasound-guided Intravascular Sonothrombolysis with a Dual Mode Ultrasound Catheter: In-vitro study

Huaiyu Wu 1, Bohua Zhang 2, Chih-Chung Huang 3, Chang Peng 4, Qifa Zhou 5, Xiaoning Jiang 6
PMCID: PMC9702596  NIHMSID: NIHMS1850194  PMID: 35201986

Abstract

Thromboembolism in vessels often leads to stroke or heart attack and even sudden death unless brought under control. Sonothrombolysis based on ultrasound contrast agents has shown promising outcome in effective treatment of thromboembolism. Intravascular sonothrombolysis transducer was reported recently for unprecedented sonothrombolysis in vitro. However, it is necessary to provide an imaging guide during thrombolysis in clinical applications for optimal treatment efficiency. In this paper, a dual mode ultrasound catheter was developed combining a 16 MHz high frequency element (imaging transducer) and a 220 kHz low frequency element (treatment transducer) for sonothrombolysis in vitro. The treatment transducer was designed with a 20 layer PZT-5A stack with aperture size of 1.2 × 1.2 mm2 and the imaging transducer with the aperture size of 1.2 × 1.2 mm2 was attached in front of the treatment transducer. Both transducers were assembled into a customized two-lumen 10 Fr catheter. In vitro experiment was carried out using bovine blood clot. Imaging tests were conducted showing that the backscattering signals can be obtained with a high signal-to-noise (SNR) ratio for the 16 MHz imaging transducer. Sonothrombolysis was performed successfully that the volume of clot was reduced significantly after the 30 min of treatment. The size changes of clot were observed clearly using the 16 MHz M-mode imaging during the thrombolysis. The findings suggest that the proposed ultrasound-guided intravascular sonothrombolysis can be enhanced since the position of treatment transducer can be adjusted with the target at the clot due to the imaging guide.

Keywords: Intravascular ultrasound transducer, Ultrasound imaging, Imaging guided sonothombolysis

I. Introduction

Venous thromboembolism (VTE) accounts for considerable morbidity and mortality and is one of the leading causes of death worldwide [1]. When the clot debris freely enters the arteries of lungs, the deep vein thrombosis (DVT) may cause the pulmonary embolism (PE), which can result in sudden death [2]. Various in the age of the clot, VTE affected three to six million patients every year in US with one third of the DVT turned into PE [3]. However, the treatment of the VTE is usually difficult and costly with an annual cumulative cost ranging from $4.9 to $7.5 billion for DVT and $8.5 to $19.8 billion for PE [4]. Traditional treatment includes the systemic thrombolysis using tissue plasminogen activator (tPA) and the mechanical thrombectomy. However, the former method suffers from long treatment time (over 24 hours) and high risk of bleeding associated complications [4]-[9], while the latter is in the risk of blood vessel damage and pulmonary embolism with large clot debris [10]-[12].

Recently, the ultrasound enhanced thrombolysis, or sonothrombolysis, was reported to be an alternative of the traditional treatment methods for improving the treatment efficacy and safety [13]-[15]. With the contrast agents of microbubbles (MBs) and nanodroplets (NDs), sonothrombolysis has shown advantage in treating both non-retracted and retracted clots [16]-[21]. By generating stable and inertial cavitations from either MBs and NDs, the agents can induce microstreaming and microjets and break the structure of the blood clots [19]. Meanwhile, the histotripsy or high-intensity focused ultrasound (HIFU) with high pressure output, was also reported to be efficient in dissolving clots with stiff and dense biostructure [22]-[27]. Besides, the dual-frequency ultrasound treatment was proved to be effective in generating higher cavitation and consequently improving the sonothrombolysis rate [28][29][30]. Most of the ultrasound treatment methods are performed with external transducer or arrays, which may be affected by the significant attenuation and aberration during the ultrasound propagation and are not suitable for the applications with ultrasound blockage and the acoustic window does not exist [31][32]. Moreover, skin burn and vessel damage are concerns due to poor acoustic coupling at the skin-transducer interface of the respiratory movement for external HIFU applications [33]. More recently, Intravascular sonothrombolysis has been proposed in our previous study, where a miniaturized intravascular transducer provides a high efficiency in thrombolysis in vitro [16][20][29][34][35]. Compared with external transducer, the intravascular sonothrombolysis can deliver the agents localy next to the clot. With lower excitation peak-negative pressure (PNP) for MB/ND-mediated sonothrombolysis, the ex-vivo tests showed no vessel wall damage [21]. However, it is important to adjust the distance between the treatment transducer and the clot for maximal lysis efficiency, which raised a new requirement for the imaging guided intravascular sonothrombolysis without using fluoscopy.

Both non-invasive and invasive ultrasound approaches have been used for blood clot detections in vitro and in vivo from several previous studies. For instance, quantitative ultrasound parameters were used to detect the clot formation in vitro under static and dynamic conditions through backscattering, attenuation, and sound velocity measurement [36]-[39]. Ultrasound elastic imaging was also applied for clot detection in vitro and in vivo. For instance, Huang et al. used shear wave approach for clot characterization in vitro [40] and strain imaging was used for DVT detection in vivo [41]. All above literatures demonstrated that ultrasound exhibits a good ability for clot detection non-invasively. Furthermore, using of needle transducer for clot and other tissues detection was performed invasively in several previous studies. For example, a high frequency needle transducer was made for cataract detection, which the needle transducer was inserted into lens directly for animal experiments [42]. For clot and plaque detection, Shih et al. designed a dual-elements needle transducer for intravascular applications through shear wave imaging [43]. All above literatures inspired us to combine the imaging transducer and treatment transducer on a new dual model ultrasound catheter for imaging guided intravascular sonothrombolysis.

In the present work, a novel design of combining imaging transducer and treatment transducer on an intravascular catheter was proposed for sonothrombolysis. The piezoelectric stacked transducer was designed and fabricated using PZT-5A as the material for active layers, with an operational frequency of 220 kHz for sonothrombolysis treatment. A 16 MHz imaging transducer was developed and assembled with the treatment transducer for imaging the sonothrombolysis process in real-time. Experiments were carried out using 3-day bovine blood clot. The ultrasound image was obtained for identifying the position of treatment transducer during sonothombolysis to demonstrate the concept of imaging and therapy with one intravascular device.

II. Materials and Methods

A. Transducer design and fabrication

Since the sub-MHz frequency of transducer exhibits a stable ability for sonothrombolysis, a center frequency of 220 KHz was used for treatment transducer in the present study [16][17]. For the single layer transducer, the small aperture size cannot provide enough pressure output because of the relative larger electrical impedance (e.g. > 1 kΩ). To reduce the electrical impedance of the transducer and increase the pressure output, a stacked transducer was designed and fabricated for treatment following the similar method reported in [16]. PZT-5A (Type III 301, TRS Technologies, Inc., State College, PA, US) was chosen as the active material due to its relatively high piezoelectric coefficients. With a thickness of 250 μm for each layer, 20 layers of PZT-5A were stacked, corresponding to the designed frequency of 220 kHz. By maintaining the pressing pressure with a customized jig, each adjacent layer was bonded with 25 μm thick conductive silver epoxy ((E-Solder 3022, Von-Roll Inc., Cleveland, OH). At the same time, 50 nm Al2O3 particles were mixed with the epoxy (EPO-TEK 301, Epoxy tech. Inc., Billerica, MA) with a volume ratio of 25% to obtain a relatively high acoustic impedance for matching. Then the stack was diced into a dimension of 1.2 mm × 1.2 mm with a dicing saw (DAD322, Disco. Inc, Boston, MA), which was small enough for fitting into the 10-Fr catheter. An Al2O3/epoxy layer was then attached to the side of the transducer as an insulation layer, which is moldable before curing process. After that, the side electrodes were connected with silver epoxy with an electric resistance less than 1 Ω. With wire connection, a 13 μm thick parylene layer was coated on the whole transducer as the passivation layer.

To detect the clot, a 16 MHz imaging transducer was designed and fabricated. A PMN-PT single crystal sample was lapped to 120 μm thick with a lapping machine (PM5 Lapper, Logitech. Ltd, San Francisco, CA). Then a layer of 200/10 nm Au/Ti was sputtered on the sample as the electrode. Alumina composite with the thickness of 120 μm was bonded as the matching layer. The backing layer was then attached and lapped to 300 μm with the conductive silver epoxy (E-Solder 3022, Von-Roll Inc., Cleveland, OH). The sample was then diced into the size of 1.2 × 1.2 mm2 by a dicing saw (DAD323, Disco Inc., Boston, MA). After the wire connection, an 8 μm thick parylene layer was coated on the whole transducer as the passivation layer. The main design parameters were shown in Table I and Table II for the treatment transducer and imaging transducer, respectively. After fabricating the two transducers, the treatment transducer was assembled into the main lumen of the catheter and fixed with epoxy (EPO-TEK 301, Epoxy tech. Inc., Billerica, MA). Then, the imaging transducer was attached in front of the stacked transducer as shown in Fig. 1(c).

TABLE I.

Design parameters for the treatment transducer

Material Thickness Velocity Density Acoustic
Impedance
Active layer PZT-5A 250 μm 4350 m/s 7750 kg/m3 33.7 MRayl
Matching layer Al2O3/epoxy 2.1 mm 2770 m/s 1300 kg/m3 3.6 MRayl
Insulation layer Al2O3/epoxy 50 μm 2770 m/s 1300 kg/m3 3.6 MRayl
Passvition layer Parylene C 13 μm 2200 m/s 1290 kg/m3 2.8 MRayl
Bonding layer E-solder 3022 (centrifuged) 25 μm 1850 m/s 3200 kg/m3 5.9 MRayl
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TABLE II.

Design parameters for the imaging transducer

Material Thickness Velocity Density Acoustic
Impedance
Active layer PMN-PT 120 μm 4600 m/s 8100 kg/m3 37.3 MRayl
Matching layer Al2O3/epoxy 700 μm 2770 m/s 1300 kg/m3 3.6 MRayl
Passvition layer Parylene C 10 μm 2200 m/s 1290 kg/m3 2.8 MRayl
Backing layer E-solder 3022 (centrifuged) 300 μm 1850 m/s 3200 kg/m3 5.9 MRayl
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Fig. 1.

Fig. 1.

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(a) Structure for the stacked transducer; (b) Schematic for the top view of the catheter; (c) Schematic for the side view of transducers and microbubble tube integrated within the catheter; (d) Photos of the fabricated catheter, including treatment and imaging transdcuers.

B. Transducer characterizations

To characterize the bandwidth of the imaging transducer, the pulse-echo test was first carried out with the 16 MHz transducer. Fig. 2 shows the setup for the pulse-echo test. A pulser/receiver (5900 PR, Olympus, WA, USA) was used to excite the transducer with a PRF of 200 Hz and pulse energy of 2 μJ with a gain of 20 dB. A steel bar was used as the reflector at a distance of 4 mm. The RF signal was collected with the oscilloscope (DSO7104B, Agilent Technologies, Santa Clara, CA, USA). The measured pulse-echo signal was used to obtain the bandwidth of the prototyped transducers.

Fig. 2.

Fig. 2.

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Schematics of the experiment setup for transducer characterizations: (a) Pulse/echo test for the 16 MHz imaging transducer; (b) Pressure output test for the 220 kHz treatment transducer.

To measure the pressure output induced by the treatment transducer, a function generator (33250A, Agilent Tech. Inc., Santa Clara, CA) was first connected to the power amplifier with a power gain of 28 dB (75A250A, AR, Souderton, PA). Then the transducer was excited with a sinusoidal pulse of 10 cycles per 10 ms in a water tank. A 20 dB preamplifier (AH-2020, ONDA Crop. Sunnyvale, CA) and a hydrophone (HGL-0085, ONDA Crop. Sunnyvale, CA) was located with a distance of 1mm from the treatment transducer surface to collect the output signal. The peak-to-peak pressure and peak negative pressure outputs were measured under the driving voltage range of 10 Vpp - 180 Vpp. An in plane (X-Y) scanning was carried out in an area of 2 × 2 mm2 to estimate the pressure output distribution, which the scanning poisiton starts at 1 mm away from the transducer suface. The focal depth and focal zone of the transducer were estimated in a range from 1 to 5 mm with a Y-Z plane scanning.

C. Blood Clot Preparation

Blood clot samples were prepared following the method in our previous works [16][20]. The acid citrate dextrose (ACD) bovine blood (Densco Marketing, Inc., Woodstock, IL, USA) was added with the 2.75% W/V calcium chloride solution (Fisher Scientific Fair Lawn NJ) with a volume ratio of 10:1. To better control the size of the clot, a PDMS channel was prepared with a 3D-printed mold. The mixture was drained into the channel and sealed. The channels were then immersed in the 37 °C water bath for 3 hours. After that, all the samples were stored under 4 °C for 3 days for retraction and the clot was formed with a length of 22 cm and diameter of 7.5 mm [44].

D. Sonothrombolysis with m-mode imaging

For the in-vitro thrombolysis test, the microbubbles reported in the previous work [16] were injected as the agents with a concentration of 108/mL at a flow rate of 0.1 mL/min. A function generator (AFG 3000, Tektronix, Beaverton, OR) was first connected to a power amplifier (75A250A, AR, Souderton, PA). A pulse signal was generated with a duty cycle of 5% with a PRF of 200 Hz (an input voltage of 80 V) on the treatment transducer, corresponding to a peak-negative pressure (PNP) of 2.8 MPa, which was considered to be high enough for the thrombolysis without obvious temperature increase for possible thermal damage [16][20][21]. In every minute, the transducer was moved forward by 0.25 mm with a 3D positioner (Anet A8, Anet Inc, Shenzhen, China), which corresponds to a treatment of 7.5 mm in 30 min. For the detection of the clot, the response from the clot was first measured in the water tank with a setup shown in Fig. 3(a). Then, the clot was detected in the PDMS tube for the tracking of the clot during the treatment. The imaging transducer was set to be stable during the treatment and after every minute treatment, a saline flushing was conducted and then the clot position was detected with the 16 MHz imaging transducer.

Fig. 3.

Fig. 3.

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(a) Experimentals setups for the (a) clot detection in water and (b) the sonothrombolysis with clot detection.

The M-mode imaging was conducted for guiding the user to move the transducer close to the clot for a better treatment during the sonothrombolysis. The envelops of the backscattering signals were obtained via Hilbert transform, then the M-mode image was obtained in a dynamic range of 40 dB. The x-axis in M-mode image represents the treatment time and the y-axis shows the distance between transducer and the surface of clot. The setup was shown in Fig. 3(b).

III. Results

A. Transducer characterizations

Fig.4 shows the measured impedance of the transducer and the pulse-echo result for the imaging transducer. Under 1 μJ input, the transducer echo reached a peak-to-peak amplitude of 1.3 V with a distance of 4 mm from the target. The −6 dB bandwidth was 54.3% with a center frequency of 16.2 MHz. The thin backing layer of the imaging transducer was applied to prevent its blockage on the treatment transducer. Fig 4(b) showed the ringdown signal due to the thin backing which decreased the center frequency and the bandwidth of the imaging transducer. Fig. 5(a) showed the measured electrical impedance curve for the treatment transducer. The working frequency was 239 kHz with an impedance of 71 Ω. The capacitance and dielectric loss were 3.07 nF and 13.18 mU, respectively.Fig. 5(b) shows the pressure output of the transducer measured using an hydrophone under the input voltage from 10V to 130V. The peak-to-peak pressure reached 5.6 MPa with 80 V input, corresponding to a peak negative pressure of 2.8 MPa, which is considered to be high enough for generating cavitation with microbubbles. Fig. 5(c) and 5(d) shows the pressure output distribution of the transducer at the focal point and in the axial direction. A convex passivation layer was added on the surface of the transducer to reach a relative deeper focal depth for the stacked transducer. However, the fabrication process also leaded an unevenness on the surface of the transducer, which may cause the grating patterns in the beam profile, as shown in Fig. 5(c). However, since the transducer covered the treatment region required in the test, the variation in the pressure output distribution did not influence the performance of sonothrombolysis in the present study. In an area of 2 × 2 mm2, the peak-negative pressure output was over 0.9 MPa on the surface of the clot, which was considered to be high enough for generating cavitation of microbubbles. Since the transducer still had a small focal depth, the distance between the clot surface and the transducer was precisely controlled during the treatment.

Fig. 4.

Fig. 4.

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(a) Measured impedance of the imaging transducer and (b) Measured pulse-echo response for the imaging transducer.

Fig. 5.

Fig. 5.

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(a) Measured electrical impedance curve of the treatment transducer; (b) Measured pressure output of the treatment transducer at 1 mm; Measured pressure output distributions of the treatment transducer: (c) from the X-Y plane at 1mm from the transducer surface and (d) from the Y-Z plane.

B. Clot detection in water

Fig. 6 show the backscattering signals from the blood clot in a water tank (a) and PDMS tube (b), respectively, which were obtained from the imaging transducer. With different orientation of the clot, the signal in the water tank showed stronger reflection from both sides of the clot boundaries while the singal from the PDMS channel showed a long backscattering signal with more attenuation. With less effect from the surrounding environment, the clot signal in the water tank showed a signal-to-noise ratio (SNR) of 18.3 dB, while the clot signal in the PDMS channel showed a SNR of 10.6 dB. Although more noise were induced in the PDMS channel, enough SNR has been acquired for distinguishing the blood clot from the surroudings.

Fig. 6.

Fig. 6.

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Measured clot signals with the imaging transducer in (a) the water tank and (b) the PDMS channel.

C. Sonothrombolysis with m-mode imaging

Fig. 7(a) shows the changes of clot size before and after treatments for 30 min and a saline flushing. The mass reduction of the clot after the treatment was about 31%. Fig.7(b) shows the M-mode imaging that obtained from the clot signals as function of time. The total data acquisition time is also 30 min using imaging transducer. The yellow dotted line in the image represents the surface of the blood clot while the multi-pixels under the yellow line represents the backscattering signal from the clotIt is obvious to observe that the clot mass was reduced during the treatment, which showed the capability of clot detection during the treatment This is a very useful to inform the operator that the treatment transducer should go forward to keep the focal zone within the blood clot. Notably, with a step of 250 μm, the imaging transducer also showed a good axial resolution, which is able to estimate step interval during the treatment.

Fig. 7.

Fig. 7.

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(a) Photo of sonothrombolysis experiments before and after the treatment and (b) its corresponding M-imaging during 30 min treatment. The detected clot front surface was labelled with the yellow dotted line.

IV. Discussion

In this work, a 20-layer treatment transducer was developed and integrated with a imaging transducer in a dual mode catheter. Compared with our previously reported work [34], the transducer provided high peak-negative pressure (> 2.5 MPa) with small aperture size (1.2 × 1.2 mm2), which is considered to be high enough for generating inertial and stable cavitations from the microbubbles for effective sonothrombolysis. After the 30 min treatment, the lysis rate was 31% by estimating the mass reduction. Although the lysis rate was lower than our previously reported work [20][29], it does not reach the limit of the treatment transducer. The feeding speed (0.3 mm/min) was intentionally slowed for estimating the accuracy of the imaging transducer and providing easier operation sequence for the flushing process during the treatment.

Meanwhile, the imaging transducer showed high sensitivity with a miniaturized aperture size, yielding a SNR of 10.3 dB for recognizing the clot sample from the environment during the treatment. Although affected by the ringdown signal, the imaging transducer showed a −6 dB bandwidth of 54% at 16.2 MHz with a pulse width of 0.2 μs, corresponding to a beam width around 0.3 mm, which was able to provide a high axial resolution for tracking the step interval controlled by the 3D positioner. For the M-imaging, Fig. 7(b) demonstrated the variation of the clot signal during the treatment which showed stronger response from 21 to 30 min while a relatively weaker response from 11 to 20 min. The clot signal difference can be attributed to the variation of the clot surface flatness after 10 min treatment. The increase of the clot signal at the later time of the treatment period may be caused by accumulation of the clot leftovers and the hardening of the clot, which was also reported in recent work [45]. This information provided by the imaging transducer is important, which helps the determination of the clot stage and the optimization of the VTE treatment for clinical applications.

For treatment transducer, the highest intensity of ultrasonic beam is near the focal zone (1 mm), which means the lysis region should always be kept within this zone to improve the treatment efficiency. Therefore, providing the real time information about clot mass reduction is very important during sonothrombolysis since it can inform the operator to advance the treatment transducer in the vessel. Using the non-invasive imaging modalities are difficult to reach this goal. For instance, the image resolution of angiography may insufficient, and a long-time exposure of X-ray is harmful to both patients and physicians. In addition, use of traditional ultrasound scanner is also inconvenient since it is difficult to maintain the scanning cross section in the same position for a long-time operation to observe the mass reduction. Combination of the imaging transducer and treatment transducer on the same catheter overcomes this problem during intravascular sonothrombolysis operation. The mass reduction was observed clearly in real-time using M-mode imaging. Since both transducers were fixed on the catheter, the treatment region and imaging region is in alignment all the time. This design is much convenient than using non-invasive imaging modalities in the real clinical applications. It allows the operator to move the catheter forward as the M-mode shows that the focal zone is out of the treatment region. However, the imaging transducer was disturbed as the treatment was excited during experiment. Therefore, a switch operation may reduce the interference between two transducers.

In the present study, the sonothrombolysis experiments were carried out as the clot was placed in the saline, which provides a convenient observation of mass reduction, as shown in Fig 7(a). Since the backscattering signal from blood clot is stronger than it from the whole blood [36], the M-mode imaging should still work even if the clot was in real vessel condition. However, the motion correction may be needed in clinical applications because the human body trembling changes the relative position between imaging transducer and clot during sonothrombolysis. The M-mode imaging resolution can be provided when a higher frequency imaging transducer is applied on the catheter. However, the ultrasound attenuation will also be increased with the higher operational frequency in the blood [39], which may reduce the SNR of backscattering signal. Therefore, the design with a 16 MHz imaging transducer is reasoanble for ultrasound-guided intravascular sonothrombolysis. In this work, the original design was to place the imaging transducer in front of the treatment transducer. However, it was difficult to perform it manually in perfect. Therefore, a slight offset between two transducers was occurred during the fabrication in the lab. To overcome this problem, the small imaging transducer may be designed and assembled into the matching layer of the treatment transducer forming a more compact structure for intravascular applications. It also should be noted that the channel was one-side blocked, which required additional flushing process for getting rid of the clot debris after the treatment. Therefore, a flow model will be added in the future work for better mimicking the human body environment with constant temperature.

V. Conclusions

An imaging-guided intravascular sonothrombolysis was demonstrated, in-vitro, in this paper with a dual mode ultrasound catheter that consists of a stacked transducer configuration combining an imaging transducer (16 MHz) and a treatment transducer (220 kHz). In vitro blood clot experiments showed that backscattering signals with a sNR of 10 - 18 dB were obtained from the imaging transducer. After 30 min treatment, the mass reduction of the clot was about 31%. This reduction was also observed in the M-mode imaging that about 8 mm clot was removed in the tube. These findings suggest that the proposed dual mode ultrasound catheter exhibits an efficient ultrasound-guided method for intravascular sonothrombolysis.

Acknowledgment

This work was supported by NIH grants R01HL141967 and R21EB027304. The authors thank Brian Velasco for assistance in the microbubble formulation.

Biographies

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Huaiyu Wu studied mechanical engineering and received the B.Sc. and M.Sc. degrees in 2013 and 2016 at Shandong University, Jinan, China, He is currently pursuing the Ph.D. degree in mechanical engineering at the Department of Mechanical and Aerospace Engineering, North Carolina State University, Raleigh, USA. His research interests focus on ultrasound transducer design for contrast imaging and intravascular sono-thrombolysis.

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Bohua Zhang received his B.E. degree in Material Physics from the Xi’an University of Technology and M.S. degree in Mechanical Engineering from George Washington University in 2012 and 2014, respectively. He is currently a Ph.D. student in the Department of Mechanical and Aerospace Engineering at North Carolina State University. His research mainly focuses on three topics: 1. New sonothrombolysis transducers design and fabrication. 2. Sonothrombolysis with Magnetic Microbubbles Under a Rotational Magnetic Field. 3. Thrombolysis catheter motion control for a continuous lysis process monitoring.

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Chih-Chung Huang (Senior Member, IEEE) was born in Taoyuan, Taiwan. He received the B.S., M.S., and Ph.D. degrees in biomedical engineering from Chung Yuan Christian University, Chung-Li, Taiwan, in 2002, 2003, and 2007, respectively. From 2006 to 2007, he worked at the NIH Resource Center for Medical Ultrasonic Transducer Technology, University of Southern California, Los Angeles, CA, USA as a Visiting Researcher, where he was engaged in research of high-frequency ultrasound imaging and development of new acoustic methods for cataract diagnosis. In 2008, he joined the Department of Electrical Engineering, Fu Jen Catholic University, New Taipei City, Taiwan, as an Assistant Professor. In 2012, he was promoted as an Associate Professor. In 2013, he joined the Department of Biomedical Engineering, National Cheng Kung University, Tainan, Taiwan. He is currently a Professor at the Department of Biomedical Engineering. He was the Deputy Director of the Medical Device Innovation Center, the Deputy Director of the Technology Transfer and Business Incubation Center, the Strategic Planning Division Director of the Research and Services Headquarters at National Cheng Kung University His research interests include ultrasonic tissue characterization, blood flow measurement, high frequency ultrasound, and ultrasonic instrument for medical applications, etc. Dr. Huang was selected as a member of IFMBE Asia-Pacific Research Networking Fellowship as well as the ordinary member of CoS representative of Taiwanese Society of Biomedical Engineering of IFMBE affiliated Society. He was the Secretary General of Taiwanese Society of Biomedical Engineering. He is an Associate Editor of the Medical Physics, Journal of Medical and Biological Engineering, and Acoustics. He is the TPC Member of the IEEE UFFC. Currently, Dr. Huang is a visiting scholar in NC State University.

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Chang Peng is an Assistant Professor of Biomedical Engineering at ShanghaiTech University He received his Ph.D. degree in mechanical engineering from the University of Florida in 2018, where he was involved in developing ultrasonic transducers for direct-contact ultrasonic fabric drying. He was then a post-doc research scholar in Micro/Nano Engineering Laboratory, North Carolina State University from 2018-2021, working on design and fabrication ultrasonic transducers for biomedical imaging and therapy applications.

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Qifa Zhou received his Ph.D. degree from the Department of Electronic Materials and Engineering at Xi’an Jiaotong University. He was postdoc fellow at the Penn State University. He is currently a professor of Biomedical Engineering and Ophthalmology at the University of Southern California. Dr. Zhou is a fellow of the Institute of Electrical and Electronics Engineers (IEEE), the International Society for Optics and Photonics (SPIE), and the American Institute for Medical and Biological Engineering (AIMBE). He has over 15 patents and has published more than 300 peer-reviewed articles in journals including Nature Medicine, Nature Biomedical Engineering, Nature Communication and Science Advances. His research focuses on the development of piezoelectric composites and piezoelectric single crystal for high-frequency high resolution ultrasonic elastography, stimulation on retian, intravascular and photoacoustic/OCT imaging as well as energy harvesting on biomedical applications.

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Xiaoning Jiang is a Dean F. Duncan Distinguished Professor of Mechanical and Aerospace Engineering and a University Faculty Scholar at North Carolina State University. He is also an Adjunct Professor of Biomedical Engineering at North Carolina State University and the University of North Carolina, Chapel Hill, and an Adjunct Professor of Neurology at Duke University. Dr. Jiang received his BS, MS and Ph.D. degrees from Shanghai Jiaotong University (1990), Tianjin University (1992) and Tsinghua University (1997), respectively. He received his Post-doctoral training from the Nanyang Technological University (1996-1997) and the Pennsylvania State University (1997-2001). He was the Chief Scientist and Vice President at TRS Technologies, Inc. prior to joining NC State in 2009. Dr. Jiang is the author and co-author of two books, 6 book chapters, 14 issued/published US Patents, 140 peer reviewed journal papers and over 120 conference papers on piezoelectric ultrasound transducers, ultrasound for medical imaging and therapy, drug delivery, ultrasound NDT/NDE, smart materials and structures and M/NEMS. Dr. Jiang is a member of the technical program committee for a few international conferences including IEEE Ultrasonics Symposium (TPC-5), SPIE Smart Structures and NDE, ASME IMECE, IEEE NANO and IEEE NMDC. He is the NanoAcoustics Technical Committee Chair for IEEE NTC, IEEE NTC Distinguished Lecturer (2018 and 2019), an editorial board member for the journal Sensors, an associate editor for the ASME Journal of Engineering and Science in Medical Diagnostics and Therapy, and Co-Editor-in-Chief of IEEE Nanotechnology Magazine. Dr. Jiang is a Fellow of ASME and SPIE.

Footnotes

Conflict of interest

Xiaoning Jiang has financial interest in sonovascular, Inc., who licensed the intravascular sonothrombolysis technology from NC state.

Contributor Information

Huaiyu Wu, Department of Mechanical and Aerospace Engineering, North Carolina State University, Raleigh, NC 27695, USA.

Bohua Zhang, Department of Mechanical and Aerospace Engineering, North Carolina State University, Raleigh, NC 27695, USA.

Chih-Chung Huang, Department of Biomedical Engineering, National Cheng Kung University, Tainan, 70101, Taiwan..

Chang Peng, School of Biomedical Engineering, ShanghaiTech University, Shanghai 201210, China..

Qifa Zhou, Department of Biomedical Engineering, University of Southern California, Los Angeles, CA 90007, USA..

Xiaoning Jiang, Department of Mechanical and Aerospace Engineering, North Carolina State University, Raleigh, NC 27695, USA.

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