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Configurable imaging platform for super-harmonic contrast-enhanced ultrasound imaging
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Configurable imaging platform for super-harmonic contrast-enhanced ultrasound imaging
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Configurable Imaging Platform for Super-Harmonic Contrast- Enhanced Ultrasound Imaging By Yang Li A Dissertation Presented to the FACULTY OF THE USC GRADUATE SCHOOL UNIVERSITY OF SOUTHERN CALIFORNIA In Partial Fulfillment of the Requirements for the Degree DOCTOR OF PHILOSOPHY (BIOMEDICAL ENGINEERING) December 2016 2 ACKNOWLEDGEMENT First, I would like to thank my advisor Prof. K. Kirk Shung. Dr. Shung has always been a role model of research to me. Without his knowledge, encouragement and guidance during my PhD, it would be a mission impossible for me to accomplish the dissertation or to overcome the obstacle during the PhD years. I would also like to thank my co-advisor Prof. Qifa Zhou, who supported, trusted and guided me through the dissertation. Dr. Zhou gave me the opportunities to do exciting research projects and have me involved in much learnful research experience, which I valued very much for my career growth. I would also like to thank the members of my qualification and dissertation committee: Dr. Krishna Nayak, Dr. Jesse Yen, and Dr. Viktor Prasanna. Their insightful inputs really helped me refine my dissertation and solved many problems. This dissertation was made possible with the help and support from many colleagues inside and outside the NIH Resource Center for Ultrasonic Transducer Technology. I am grateful to the contribution from Dr. Fan Zheng, Dr. Jianguo Ma, Dr. K. Heath Martin, Dr. Kwoko Lam, Dr. Jun Zhang, Dr. Yao-sheng Tung, Dr. Ming Qian, Dr. Ruimin Chen, Dr. Changgeng Liu, Dr. Teng Ma, Ms. Mingyue Yu, Dr. Hojong Choi, Dr. Bonjing Kang, Dr. Yunhua Rao, Mr. Jason Ho, Mr. Hayong Jung, Mr. Harry Chiu, Mr. Xuejun Qian, Mr. Zeyu Chen, Dr. Ying Li and Dr. Andrew Weitz, as well as all other members in USC UTRC. I would also like to thank the collaborators from UNC and NCSU. I would like to thank my dearest friends, Ms. Jiayin Kuang for her consistent company, assistance and encouragement during my study and living in USC from the very beginning. Finally, this dissertation is dedicated to my parents, my grandparents and my friends for their love and unlimited support. Thank you for always being there for me. 3 TABLE OF CONTENTS ACKNOWLEDGEMENT .................................................................................................. 2 TABLE OF CONTENTS .................................................................................................... 3 LIST OF TABLES .............................................................................................................. 6 LIST OF FIGURES ............................................................................................................ 7 ABSTRACT ...................................................................................................................... 12 Chapter 1 Introduction ...................................................................................................... 14 1.1 Super-Harmonics Contrast-Enhanced Ultrasound: Acoustic Angiography ... 14 1.2 Motivation and Objective ............................................................................... 17 1.3 Outline of the Dissertation .............................................................................. 18 Chapter 2 Physics and Theory of Ultrasound Imaging ..................................................... 20 2.1 Ultrasound Wave Propagation ........................................................................ 20 2.2 Ultrasound Transducers .................................................................................. 25 2.2.1 Single Element Transducer .............................................................. 25 2.2.2 Array Transducer ............................................................................. 28 2.3 Ultrasound Imaging ........................................................................................ 29 2.3.1 Pulse Echo and A-line ...................................................................... 29 2.3.2 Beamforming for Array Imaging ..................................................... 31 2.4 Imaging System .............................................................................................. 34 Chapter 3 Principle of Super-Harmonics Imaging ............................................................ 36 3.1 Microbubble Contrast Agent Linear Modeling ............................................... 36 3.2 Microbubble Contrast Agent Nonlinear Modeling ......................................... 38 4 3.3 Principle of Super-harmonics Imaging ........................................................... 40 3.3.1 MCA Preparation ............................................................................. 41 3.3.2 Behavior Simulation ........................................................................ 42 Chapter 4 Configurable Dual-Channel System for Super-Harmonic Imaging ................. 48 4.1 General Design Consideration ........................................................................ 48 4.2 Specific System Design .................................................................................. 51 4.2.1 FPGA Logic ..................................................................................... 51 4.2.2 Analog Receiver Design .................................................................. 55 4.2.3 Pulse Generator ................................................................................ 56 4.3 Transducer Design .......................................................................................... 57 4.4 Experiment Setup ............................................................................................ 59 4.5 System Characterization ................................................................................. 60 4.5.1 Analog Receiver Characterization ................................................... 60 4.5.2 Pulser Characterization .................................................................... 62 4.6 Imaging Result ................................................................................................ 64 4.7 Discussion ....................................................................................................... 69 4.7.1 Selection of Transmit and Receive Frequency ................................ 69 4.7.2 Resolution ........................................................................................ 70 4.7.3 Dynamic Range ................................................................................ 71 4.8 System Upgradation ........................................................................................ 72 Chapter 5 Ultrasonic Array System for Super-Harmonics Imaging ................................. 74 5.1 64-channel Ultrasound Beamforming System ................................................ 74 5.1.1 Transmit Beamformer ...................................................................... 75 5.1.2 Receive Beamformer ....................................................................... 76 5.1.3 Transducer and System Calibration ................................................. 78 5.2 Sampling Rate Upgradation ............................................................................ 79 5 5.3 Single-Transmit-Array-Receive (STAR) Mode ............................................. 84 5.4 Phase Array Imaging Result ........................................................................... 86 5.4.1 Experiment Setup ............................................................................. 86 5.4.2 Imaging Result ................................................................................. 89 5.5 Linear Array Imaging Result .......................................................................... 90 5.5.1 Experiment Setup ............................................................................. 90 5.5.2 Imaging Result ................................................................................. 90 5.6 Summary and Discussion ................................................................................ 96 Chapter 6 Synthetic Aperture Beamforming for Super Super-harmonic Imaging ........... 97 6.1 Introduction to Virtual Point Synthetic Aperture Beamforming .................... 97 6.1.1 VPSA for Single Element Imaging .................................................. 97 6.1.2 VPSA for Array Imaging ............................................................... 100 6.2 Application of VPSA in Bmode Imaging ..................................................... 102 6.2.1 Application of VPSA to Ultrasound Bio-microscopy ................... 102 6.2.2 Application of VPSA to IVUS Imaging ........................................ 106 6.3 Application of VPSA for Super-Harmonic Imaging .................................... 108 6.3.1 VPSA for Dual-Element Super-Harmonic Imaging ...................... 109 6.3.2 VPSA for STAR Mode Super-Harmonic Imaging ........................ 114 6.4 Summary and Discussion .............................................................................. 117 Chapter 7 Conclusion and Future Work ......................................................................... 118 7.1 Conclusion .................................................................................................... 118 7.2 Future Work .................................................................................................. 119 7.2.1 Dual-frequency Array with Dynamic Transmit/Receive ............... 119 7.2.2 In-vivo Animal Study ..................................................................... 120 BIBLIOGRAPHY ........................................................................................................... 121 6 LIST OF TABLES Table 4-1 Devices Used in the Proposed System ............................................................. 51 Table 4-2 Major User Defined Parameters of the Proposed System ................................ 53 Table 4-3: Transducer Used For Single-element Super-harmonic Imaging ..................... 58 Table 5-1: Specifications of the 30 MHz prototype linear phased array [21] .................. 78 Table 5-2: Measurement of axial and lateral resolution of linear array STAR mode super-harmonic imaging at different transmit voltages. ................................................... 94 Table 6-1: Transducer Used for Super-harmonic Beamforming Study .......................... 109 Table 6-2: Resolution Improvement B-mode VPSA+CFW ........................................... 111 Table 6-3: Resolution Improvement for Linear Array Super-harmonic STAR Mode Imaging VPSA+CFW ..................................................................................................... 117 7 LIST OF FIGURES Figure 1-1: A comparison of maximum intensity projections of image data for the same kidney acquired with two different modalities: (A) acoustic angiography. (B) contrast enhanced CT. Scale bars in both images indicate 1 cm. A cartoon illustrates the approximate boundaries and orientations of these images (C). [11] ........................... 16 Figure 2-1: Reflection and refraction of acoustic wave. [16] ........................................... 21 Figure 2-2: Attenuation coefficient of various tissues and biologically relevant liquid as a function of frequency. [17] ........................................................................................ 22 Figure 2-3: In-vivo backscatter coefficient of dermis and of other tissue. [19] ................ 24 Figure 2-4: Structure of a single element transducer. ....................................................... 26 Figure 2-5: Structure of array transduce and the dimension system. [20] ........................ 28 Figure 2-6: Basic principle of the ultrasound pulse-echo and the acquisition of an A- line [16]. ............................................................................................................................ 30 Figure 2-7: Linear scan and phase steering scan. (A) Electronic focusing by applying delay. (B) Linear scan. (C) Electronic steering and focusing by applying delay. (D) Phase steering scan. [21] ................................................................................................... 33 Figure 2-8: Receive beamforming on a phase array. [21] ................................................ 34 Figure 2-9: Block diagram of a digitized ultrasound Duplex image system, with power and spectral Doppler function. For single element transducer mechanical scanning, only the components circled in the red are necessary. [22] .............................. 35 Figure 3-1 (A) Calculation of scattering cross-section of a gas bubble, as function of transmit frequency and different radius. (B) Experiment data and theoretical curve of attenuation coefficient as a function of frequency for Albunex. [24] ............................... 38 Figure 3-2 The change in diameter (left) and in scattering-cross section (right) of free-bubble(up) and Albunex. Acoustic driving frequency of 2MHz at the pressure amplitude of 50kPa is given. [24] ..................................................................................... 39 Figure 3-3 1 st and 2 nd scattering cross-section as a function of driving frequency, left: free gas bubble; Right: Albunex. [24] ............................................................................... 40 Figure 3-4: Left: Vials of lipid and gas before and after mixing. The white froth is the suspension of microbubbles. Middle: view of microbubble under the microscope. Right: A cartoon shows the structure of the bubble and its lipid shell. [11] ..................... 42 Figure 3-5: The transfer function of microbubble with shell of different diameter. ......... 44 8 Figure 3-6: Driving pulse and the bubble radius oscillation. 0.8µm diameter bubble. ..... 45 Figure 3-7: The spectrum of scattered pressure of bubble with different diameters. The x-axis is the frequency from 0-40Mhz, y-axis is the relative amplitude to the driving pulse train, from -120 to -180dB. The blue curves are the driving pulse spectra for reference. ......................................................................................................... 46 Figure 3-8: Left: bubble diameter distribution; Right: The mean scattering cross- section spectrum of the bubble population the magnitude is normalized to dB relative to the 5MHz component. ................................................................................................... 47 Figure 4-1: Up. Block diagram of the configurable dual-frequency super-harmonic imaging system. Different color shows different power domain: blue: digital supply; green: high voltage for pulser; red, analog supply; purple, low-noise analog DC supply for ADCs. Bottom: Photography of the system. ................................................... 50 Figure 4-2: (A) FPGA logic block diagram. Dashed lines show the clock signal. (B) Logic timing of one frame data acquisition. Pulse number 1 and 3 are bipolar pulses for B-mode imaging; Pulse number 2 and 4 are multi-cycle pulse-train for super- harmonics imaging. ........................................................................................................... 54 Figure 4-3: (A) LabVIEW data acquisition software block diagram. Dashed lines indicated the concurrent stages. (B) GUI for raw data acquisition: 2000line/frame, 100Hz PRF, online software filtering and pulse echo mode. ........................................... 55 Figure 4-4: (A) Acoustic stack of the dual-layer transducer. (B) Photo of the transducer front looking. The low frequency element is the 4mm dark color block. The high-frequency element is the light color piece at the center. (C) Experiment setup for vasa vasorum phantom imaging. (D) Microbubble size distribution. ................ 58 Figure 4-5: Impulse response: tested by applying a short pulse from a pulse generator Panametrics 5900PR to the transducer connector, and collect the digitized signal from the proposed system. (A) Input pulse voltage (V) collected by the 2 GHz digital oscilloscope (dashed black) and the received signal (scaled to 1, blue) collected by the system. (B) The spectrum of the input pulse (dashed black), of the received signal (blue) and of analog filter simulation (red). ........................................................... 61 Figure 4-6: (A) Broad-band pulse and spectrum measured by oscilloscope. (B) 6MHz pulse train and spectrum measured by oscilloscope (C) Pulse echo signal of the 35MHz element collected by the proposed system, shown in normalized amplitude. (D) Acoustic pressure output of the 6MHz element measured by the hydrophone, shown in hydrophone voltage. ..................................................................... 63 Figure 4-7: Acoustic pressure collected by the hydrophone, calibrated to pressure. Measured with 6MHz 1-cycle or 2-cycle transmit pulse. ................................................. 64 Figure 4-8: In vitro B-mode image of the rabbit coronary artery obtained by 35MHz transducer element. ........................................................................................................... 64 9 Figure 4-9: Phantom acoustic angiography image. (A) B-mode image, when water is injected into the micro tube. The location of tube is circled with red. (B) Super- harmonic image obtained when MCAs are injected into the micro-tube. ........................ 67 Figure 4-10: Expanded view of super-harmonic image. (A) is redraw from Figure 4-9(B). (B) and (C) are collected using a different transducer with narrower bandwidth. (C) is collected where the micro-tube runs parallel to the imaging plane. 200um along the axial direction is indicated by the red bars. ........................................... 68 Figure 4-11: Photograph of the second generation of dual-channel system. .................... 73 Figure 5-1: Overall high-frequency ultrasonic array system architecture. [21] ................ 74 Figure 5-2: (A) Photograph of the developed ultrasonic phased array system. (B) Transmit beamformer. (C) Receive beamformer. [21] ..................................................... 75 Figure 5-3: Block diagram of the 64-channel transmit beamformer. ............................... 76 Figure 5-4: Logic Block diagram of the 64-channel receive beamformer. [21] ............... 77 Figure 5-5: (A) 30MHz phase array. (B) US image of a wire phantom by a 30 MHz PMN-PT kerfless phased array. [21] ................................................................................ 78 Figure 5-6: Transmit delay generation algorithm. Setup violation occurs at the Ө sign. ........................................................................................................................................... 81 Figure 5-7: Modified transmit delay generation algorithm. .............................................. 82 Figure 5-8: Delay computation algorithm. ........................................................................ 82 Figure 5-9: (A) A cartoon shows single-transmit-array-receive (STAR) super- harmonic imaging. The transmit beam approximation is shown as red dash line. ........... 85 Figure 5-10: (A) Simulation of the 3.5MHz single element transducer beam-profile used for phase array super-harmonic imaging. (B) Setup of the experiment, where the single element transduce was put on the side of the phase array. The microbubble filled tube is covered by the focal point of the transmit transducer. (C) Photograph of the phase array system based STAR mode super harmonic imaging. The zoom-in view of the circled region is show on the left. .................................................................. 87 Figure 5-11: (A) B-mode image of the vessel mimicking tube acquired by the phase array while the 3.5MHz single element transducer transmitting. (B) Super-harmonic signal, where excitation signal of the array was turned off. 25dB dynamic range, using two-way distance calculation. (C) Super-harmonic image acquired at 6mm, one-way distance. (D) Super-harmonic image acquired at 8mm, one-way distance. 20dB dynamic range. ........................................................................................................ 88 10 Figure 5-12: (A) Simulation of the 5.8MHz single element transducer beam-profile used for linear array super-harmonic imaging. (B) Setup of the experiment, where the single element transduce was put on the side of the linear array. The microbubble filled tube is covered by the focal point of the transmit transducer. ................................. 91 Figure 5-13: Super-harmonic image collected using the linear array. (A)- (E): transmit voltage 22, 38, 54, 70, 85 V pp respectively. 25dB dynamic range. (F) pressure output vs. transmit voltage for the 5.8 MHz single element transducer. ............ 92 Figure 5-14: Mechanism for pseudo-enhancement. .......................................................... 92 Figure 5-15: (A) Axial resolution of super-harmonic image at different transmit voltages. (B) Lateral resolution of super-harmonic image at different transmit voltages. ............................................................................................................................ 93 Figure 5-16: Super-harmonic image acquired before and after the microbubble stop pumping through the micro-tube. (A)- (H) consecutive image collected with 0.5s interval. The pump was turned off at the time of (B). 25dB dynamic range. ................... 95 Figure 6-1: (A) Concept of traditional mechanical scan B-mode imaging, where the acoustic field of the transducer is showing in dashed line. The red dot denote the transmit focus. The green dots denote the imaging point. (B) Concept of virtual point SA imaging. The number in the image shows the number of overlapping acoustic field. ................................................................................................................................ 100 Figure 6-2: Illustration of coherent factor weighting. Left: cases where the delayed synthetic beams are perfectly coherent; Right: case where the synthetic beam are not well coherent. The vertical axis is the depth, and the horizontal axis is the synthesized aperture direction. [56] ................................................................................ 100 Figure 6-3: Virtual source concept for linear array beamforming. (A) Traditional transmit focusing; (B) spherical defocusing wave transmit used for SA. [58] ............... 101 Figure 6-4: (A) Simulation of the 64MHz transducer beam-profile. (B) B-mode cross-section image of the rat aorta tissue. ..................................................................... 103 Figure 6-5: (A) Linear weighting function. (B) CFW. ................................................... 104 Figure 6-6: VPSA result: (A) VPSA with linear weighting. (B) VPSA with CFW. (C) compounding image using pre-beamforming images acquired at different depth. 50dB dynamic range. ...................................................................................................... 105 Figure 6-7: (A) Pre-beamforming IVUS image of rat aorta. (B) VPSA beamformed image of rat aorta. (C) Pre-beamforming IVUS image of size 20µm wire targets at five different depths.(D) VPSA beamformed image of wire targets. (E) Zoom-in view of (C). (F) Zoom-in view of (D). ............................................................................ 107 11 Figure 6-8: (A) Photograph of the dual-frequency transducer. (B) B-mode imaging of wire phantom using the 30MHz element. (C) VPSA+CFW beamforming result. 40dB dynamic range. ...................................................................................................... 110 Figure 6-9: Signal spectrum acquired using 5.5MHz transmitting 30MHz receiving at different transmit voltages. (A) air tube. (B) water tube. (C) MCA-filled tube. ............ 111 Figure 6-10: (A) 30MHz B-mode imaging of the microtube in the tissue phantom. (B) Pre-beamformed super-harmonic image. (C) VPSA+CFW beamformed image. (D) Longitudinal view pre-beamformed super-harmonic image. (E) Longitudinal view VPSA+CFW beamformed image. 40dB dynamic range. ............................................... 113 Figure 6-11: Result of VPSA+CFW method for phase array STAR mode super- harmonic imaging. (A) DAS beamforming, reproduced from Figure 5-11(C). (B) DAS+VPSA beamforming. ............................................................................................ 115 Figure 6-12: Result of VPSA+CFW method for linear array STAR mode super- harmonic imaging shown in Figure 5-13. (A)- (E) transmit voltage 22, 38, 54, 70, 85 V pp respectively. 25dB dynamic range. ........................................................................... 116 12 ABSTRACT Super-harmonic contrast-enhanced ultrasound has been demonstrated to be useful for microvascular imaging. By exciting the microbubble contrast agent near its resonance frequency and receiving echo at the super-harmonic, high-contrast high-resolution image from contrast agent could be obtained and discriminated from tissue. While dedicated dual-element transducers have been designed, no existing integrated system could fully support the application. To fulfill this need, a configurable dual-channel transmit/receive system for super-harmonic imaging was designed, built and tested experimentally. The system was capable of obtaining both B-mode imaging and super-harmonic imaging. Multiple user-defined working modes were selectable, implemented by a field programmable logic gate. The function of pulse generator, low- noise data acquisition unit and configurable signal processing unit were integrated in the system. The system had 56dB SNR and 42dB variable gain. When working with a 6.5MHz, 35MHz dual-frequency transducer, the system was capable of obtaining super- harmonic contrast images of 30dB dynamic range on a vasa vasorum mimicking intravascular phantom. The system showed promise for microvascular assessment in preclinical studies. Currently, all super-harmonic imaging studies were based on dual-element transducer. Array transducer based imaging, which provides transmit/receive from multiple channels synchronously, could potentially improve the signal coherence and imaging resolution, as well as enhance the speed of imaging. Based on an array beamforming system, with proper software and hardware upgradation, we successfully demonstrated the capability of single-transmit-array-receive mode (STAR) super- 13 harmonics imaging. The system was capable of working with multiple kinds of transducer for the application. The beamforming method for super-harmonic imaging was studied. Synthetic aperture beamforming, which was a well-accepted method in traditional ultrasound, was used for super-harmonic imaging for the first time. By utilizing the echo information from the superimposed acoustic field of the neighboring emission, an image with higher lateral resolution, higher SNR could be acquired. We demonstrated the method in B- mode ultrasound microscopy, and showed that the method works for super-harmonic imaging as well, in both dual-element and array approaches. 14 Chapter 1 Introduction 1.1 Super-Harmonics Contrast-Enhanced Ultrasound: Acoustic Angiography Angiography was the medical imaging techniques to visualize arteries, veins, heart chambers and other vessel structure of the body. It was used to prescreen and to diagnose the state and progression of many diseases, such as atherosclerosis, which was the most common cause of cardiovascular diseases. Angiography was first developed based on X-ray imaging. Nowadays, because of the using of digital imaging system, the method was often referred to Digital Subtraction Angiography (DSA), where the vessel image could be obtained by applying radio-opaque contrast agent, and by subtracting the background organs information from the contrast-enhanced image. While X-ray contrast angiography still remained the gold standard, other imaging modalities were also developed and provided to be useful for vascular imaging, such as computer tomography (CT), magnetic resonance imaging (MRI) and Duplex ultrasonography (B-mode ultrasound imaging + Doppler flow detection [1-3]. Duplex ultrasound has higher temporal resolution compared to the other imaging modalities mentioned above. It was also non-radioactive, relatively low-cost and light weight, making it the mainstay routine testing for vascular assessment[4]. Because blood was a weak scatter of acoustic wave, B-mode ultrasound could not be used to image the blood vessel. Doppler ultrasound inferred the flow information from the frequency shift, which provided the flow velocity and direction information beside the locational image. 15 Contrast-enhanced ultrasound was another ultrasonic imaging method visualized blood vessel and tissue with weak acoustic scattering by applying contrast medium such as shell-encapsulated air bubbles [5, 6]. These microbubble contrast agents (MCAs) were much more compressible compared to tissue, thus could significantly enhance the acoustic backscatter. MCAs were used as a contrast enhancer during B-mode or Doppler ultrasound. By binding certain ligand identifiers to the shell of MCA, it could also preferentially bind to endovascular markers of interest, selectively identifying regions of tissue with specific functions [7], which provided functional information on the molecular level beside the anatomical structure. Moreover, by making use of the high- frequency components generated from the MCAs nonlinear oscillation, imaging resolution could be enhanced compared to other sonography angiography methods. This new method is called super-harmonics contrast-enhanced ultrasound imaging (for simplify super-harmonics imaging). Super-harmonic imaging utilized the high-frequency signal generated from the nonlinear behavior of MCAs under acoustic pressure to provide high-resolution angiography images. As shown in Figure 1-1, the super-harmonic image of the rat right kidney microvasculature was obtained with a comparison of CT imaging [8], where a resolution of 100µm was achieved. It has also been demonstrated that super-harmonic imaging was capable of distinguishing between the healthy and the tumor-bearing tissue in a rodent model, by providing high-spatial-resolution images for the angiogenic processes associated with tumor development [9]. Another potential major application of super-harmonic imaging was for coronary artery assessment when integrated with catheter-based intravascular ultrasound (IVUS) technology [10]. IVUS utilized miniature 16 transducer with 10 to 40MHz center frequency to imaging the lumen of blood vessel through endoscopic procedure. The thickness of vessels, plaque size, morphology and composition could be obtained using IVUS. When dual-element dual-frequency IVSU transducer was used, super-harmonics imaging could be applied to image the supporting micro-vessel networks that provide nourishment and oxygen to the outer component of the vessel wall, namely the vasa vasorum, as shown in the phantom studies from [10]. Though many of the studies were still in the early stage, super-harmonic imaging showed great promises for micro-vasculature angiography in clinical and preclinical applications. Figure 1-1: A comparison of maximum intensity projections of image data for the same kidney acquired with two different modalities: (A) acoustic angiography. (B) contrast enhanced CT. Scale bars in both images indicate 1 cm. A cartoon illustrates the approximate boundaries and orientations of these images (C). [11] 17 1.2 Motivation and Objective On the imaging system side, super-harmonics imaging required a transmit pulser with adjustable voltage and multi-cycle output to take care of the variance among transducers transmission efficiency, especially for the IVUS transducer with very small aperture size. In addition to contrast-enhance image, high-frequency B-mode image of the tissue were necessary during acoustic angiography structure reference, which was acquired by doing the pulse-echo of the high-frequency transducer. Previous study used commercial system (Vevo 700, VisualSonics, Toronto, Canada) for pulse-echo imaging and harmonic signal receiving, and used independent power amplifier and function generator for excitation pulse generation [12]. To reduce the system complexity, to improve the noise control and to enhance the portability for super-harmonic imaging pre- clinical application, an integrated imaging system for the application was needed. A few research imaging systems have been developed for high-frequency IVUS or dual-channel application [13-15]. But to the best of the authors’ knowledge, none of the lab-made or commercial imaging systems were fully capable of super-harmonic imaging. So far, all the super-harmonic imaging was based on the mechanical scan, where integrated dual-element transducer, either for transcutaneous or IVUS application, was mounted on a motorized scanning system, and B-mode images were acquired by compounding A-lines from individual scans. The speed of imaging was thus very slow, which becomes a problem for imaging of moving tissue, such as in the application of cardiovascular studies. Using array transducer with multiple elements provided a solution for the imaging speed, which required the usage of multi-channel electronics system to handle the transmitting, receiving and processing to form an image, namely a 18 beamformer. Along with the development of digital electronics technology, array transducer and digitized beamformer system with massive computation ability became the trend of all sonography imaging. However, currently there is no array-based study in the area of super-harmonics imaging. The motivation of this study was to investigate the electronically hardware requirement for super-harmonic imaging, to study how to use array transducer for super- harmonics imaging, and to build a beamforming method for the application. A dual- channel transmit/receive system for super-harmonics imaging was designed, to study the hardware capability of the imaging system for super-harmonic imaging. A field- programmable-logic-array (FPGA) based beamformer system would be applied for the array-based super-harmonic imaging. Beamforming method for super-harmonic imaging would be investigated and implement. Results from phantom and in vitro imaging experiment would be shown. 1.3 Outline of the Dissertation The rest of this dissertation was organized as follows: In Chapter 2: basic theory of ultrasound propagation, ultrasound transducer and high-frequency ultrasound imaging. In Chapter 3: theory and physical of microbubble contrast agent, behavior simulation of microbubble under super-harmonic imaging condition. In Chapter 4: the design, fabrication and experiment result of a configurable dual- channel system for super-harmonic imaging. 19 In Chapter 5: introduction of the upgradation a 64-channel array system, single- element-transmit-array-receive super-harmonic imaging. In Chapter 6: virtual point synthetic aperture beamforming and coherent factor weighting, application of the beamforming method to ultrasound microscopy, dual- element super-harmonic and single-element-transmit-array-receive super-harmonic imaging. In Chapter 7: conclusion of the dissertation and future work. 20 Chapter 2 Physics and Theory of Ultrasound Imaging In this chapter, the physics of wave propagation and basic ultrasound imaging will be introduced. Design factor of ultrasound transducer will be discussed, with an introduction to some critical imaging criteria and a brief introduction about imaging system. 2.1 Ultrasound Wave Propagation Ultrasound wave is a high-frequency mechanical wave with the frequency from 20 kHz to Gigahertz range. Propagation of ultrasonic wave can be described by the wave equation [16]: !" ## !$ + !" #& !' + !" #( !) = 𝜌 ! , - !. , (2-1) where 𝜅 $$ ,𝜅 '$ 𝜅 $) are the longitudinal stress in Z (axial) direction, and the shear stress in the two lateral directions, respectively, 𝜌 is the mass density and w is strain in the axial direction. A few physical processes affect the propagation of ultrasound through tissue: -- Reflection and refraction occur when the wave encounters an interface between two media with different acoustic impedance; see Figure 2-1, similar to what happened to the light, given by the equations: 1 2 1 3 = 4 , 5678 3 94 : 5678 ; 4 , 5678 3 <4 : 5678 ; = (2-2) 21 1 ; 1 3 = >4 : 4 , 5678 3 (4 , 5678 3 <4 : 5678 ; ) , (2-3) where Z i is the acoustic impedance. Ii, Ir and It are the incident, reflected and transmitted acoustic intensity, defined by: 𝐼 = B = C , D5 = B = C , 4 (2-4) where p is peak value of the pressure, c is the sound velocity in the medium. Figure 2-1: Reflection and refraction of acoustic wave. [16] -- Attenuation: for a plane wave propagating in the Z-direction, given the attenuation coefficient α, the pressure decreases exponentially denoted by the equation: p z =p G e 9IJ (2-5) α is made up of two term, the absorption coefficient α L and the scattering coefficient α 7 . Attenuation coefficient is frequency dependent and has strong dependence on the type of tissues. It can be defined as: 22 α=α G f N (2-6) where 0 is the attenuation coefficient at 1 MHz and γ is the frequency dependence parameter. Attenuation vs. frequency function of various tissue is shown in Figure 2-2. -- Absorption: absorption is a loss of ultrasound energy due to viscosity and relaxation process of the medium. It can be expressed in the format of Equation 2-6. In some tissue such as liver, the absorption is dominated attenuation effect compared to scattering. Figure 2-2: Attenuation coefficient of various tissues and biologically relevant liquid as a function of frequency. [17] 23 -- Scattering: as the wave incident upon a particle, part of the wave will be scattered to all direction, acting as a point source acoustic wave, which can be described by a term called differential scattering cross-section: σ P (o,i)= 1 S T , 1 3 (2-7) where Ir and It are incident and scattered acoustic intensity. R is the distance from the observer to the scatter. σ P is a function of the incident (i) and observer (o) direction. When o=-i, it is called the backscattering cross-section. It reflects the power scattered by the particles that encompasses an area of R 2 . Integrating the different scattering cross- section by solid angle 4π gives the scattering cross-section: σ 7 = σ P 𝑑𝛺 = 4𝜋𝑅 = 1 S 1 3 4π (2-8) Analytical solution of σ 7 can be calculated [18]. For example, the scattering cross- section for a red blood cell is 1.1x10 -12 cm 2 . For 1 W/cm 2 intensity, only 1.1x10 -12 W of power is scattered by the cell [16]. In a dense distribution of scatters, such as in bio-tissue, the scatter-to-scatter interaction cannot be ignored. Volumetric backscattering cross-section was often used in such as cas, also sometimes called the backscattering coefficient. The physical meaning of backscattering coefficient is the scattering in a unit volume, with a unit of cm -1 sr -1 , where sr is the unit of solid angle. The backscattering coefficient in blood and dermi tissue has been measured in previous studies[19] , some of the results are shown in Figure 2-3. 24 Scatter cross-section is an important measurement when consider the MCAs behavior. We will discuss it in Chapter 3. Figure 2-3: In-vivo backscatter coefficient of dermis and of other tissue. [19] -- Nonlinear Parameter B/A: the nonlinear behavior of a fluid medium can be expressed by a second-order parameter B/A. A and B are the first and second order coefficient in the Taylor series expansion of the function of pressure wave: 𝑝 =𝑝 𝑠 G ,𝜌 G +𝐴 D9D _ D _ + B = 𝐵 D9D _ D _ = +⋯ (2-9) where 𝑠 G and 𝜌 G are equilibrium entropy and density. Second harmonic signal from tissue can be used for imaging. However, the high order harmonic from the tissue was too weak to be detected. 25 2.2 Ultrasound Transducers 2.2.1 Single Element Transducer The key component of ultrasound imaging is the piezoelectric transducer, which converts electrical energy to mechanical energy or vice versa. When electrical field is applied to the piezoelectric material, realigning of the dipoles in the material results in a deformation of the material thickness [16], which transfers to acoustic or ultrasonic wave. This effect is called piezoelectric effect. In contrast, reverse piezoelectric effect is the transferring from mechanical wave to changing electrical filed. Ultrasound imaging utilizes both piezoelectric and reverse piezoelectric to transmit and to receive the ultrasound wave. Single-element transducer is the elementary type of ultrasonic transducer. A broadband single-element transducer normally consists of a piezo-element which is sputtered with electrodes on both surfaces, matching layers, backing material and sometimes a lens, shown in Figure 2-4. The resonance frequency, or the center frequency, of the transducer is determined by the thickness of the transducer, given by: 𝑓 G = c5 =d (2-10) where L is the thickness of the material and the n is an odd number. The bandwidth of the transducer is determined by the matching and backing layer property. Matching layer is inserted between piezoelectric materials and the loading medium to compensate the acoustic impedance mismatch, thus most of energy could be transmitted to the front size. According to transmission line theory, transmission has no loss when thickness of the 26 matching material is equal to 4 m ( m was the wavelength of the matching material) and acoustic impedance of the matching material m Z satisfied: 𝑍 B = 𝑍 f 𝑍 G B/= (2-11) which can be derived from Equation 2-3. For practice, 1/3 exponential factor is often used, and two layers of matching are necessary to improve the transmit coupling [16]. Similarly, when the pressure wave moves backward and hits the rear surface of the piezo-element at a normal incidence, part of the energy is reflected back and the rest is sent into the backing layer. The reflected wave causes ringing in the produced acoustic signal, thus impairs the bandwidth. However, sensitivity of the signal will reduce if Z2 is selected close to Z0, in which case there is no ringing effect. A compromise has to be made between bandwidth and sensitivity in practical applications. Figure 2-4: Structure of a single element transducer. 27 For single element transducer with circular aperture, if flat surface is used, the natural beam focus location is given by: 𝑧 G = i , j (2-12) where 𝑎 is the radius of transducer and λ is the wavelength. z 0 separates the acoustic field into near-field and far field. For imaging, target of interest should be always in the far field or beyond z 0. The analytical far-field beam-profile function can be solved in the far field, given which we can calculate the size of the main lobe of a circular aperture transducer: sin𝜑 =0.61 𝜆 i (2-13) where 𝜑 is the angular direction. Acoustic lens or press-focused convex aperture could be applied to focus the beam closer. Given a focal distance of z, a factor caused “f number” was defined by: 𝑓 # = $ =i (2-14) 𝑓 # is an important parameter from which we could empirically estimate the beam diameter at the focal point (W b ) and the -3dB intensity region depth (depth of focus, D f ): 𝑊 v ≈𝑓 # 𝜆 (2-15) 𝐷 y ≈𝑓 # = 𝜆 (2-16) 28 2.2.2 Array Transducer Array transducer is the trend for ultrasonic imaging. Array employs many piezoelectric elements, where each element has its own electrical connection. Elements can be excited individually or in groups, effectively forming a sub-aperture. Like a single-element transducer, a backing material and one or two matching layers are used to improve arrays’ performance. To minimize acoustic cross-talk, piezoelectric material is diced into small elements. The space between two elements is called kerf. The distant between individual elements is called pitch. The kerfs might be filled with acoustic isolating material. The isolating material serves a dual purpose: it can decrease acoustic cross-talk, but it also provid a rigid support for array elements. Sometimes, a lens is used to focus the acoustic beam in the elevation direction. Figure 2-5: Structure of array transduce and the dimension system. [20] The radiation beam-profile of an array in the far field is given by: 29 ( ) 1 sin sin N m bu Lu Hu c u m c g = = (2-17) where sin x u= , ( ) Hu is the direction function at an angle of x , L is length of the array, N is the number of the elements, b is the width of the element, g is pitch, is the wavelength in the loading medium. * indicated the operation of convolution. Sinc is the sinc function: sinc(x)=sin(x)/x. High side lobes called grating lobes appear at certain angles because interference, which are related to the wavelength and the pitch by the following equation: 𝜑 z =sin 9B ( cj z ) (2-18) where n is an integer. From the above equations, we could also conclude that: the width of the element should be as large as possible in order to damping the magnitude of the grating lobes; the aperture of the array should be as large as possible in order to get a narrower main beam width. Em pirically, pitch of the array is usually selected between 0.5λ and 2λ, thickness of the element ( b t ) should be < 0.6 to avoid lateral resonance [16]. However, above requirements could not be satisfied at the same time, some trade-offs have to be made in a practical design. 2.3 Ultrasound Imaging 2.3.1 Pulse Echo and A-line When electric pulse is delivered to the transducer, the wave fronts of the ultrasound wave generate and propagate through the tested object. When the wave fronts 30 reach the discontinuities of acoustic impedance through the propagation, partially of the wave is scattered back, and received by the same transducer. The distance of different objected of discontinuity could be inferred from the time difference of the receiving signal following the equation: 𝑑 = .∙5 = (2-19) where c is the speed of sound in the test object. This is called pulser-echo mode ultrasound, and the signal acquired is called an A-line, see Error! Reference source not found.. If converting the amplitude of the echo to pixel in gray level, and doing multiple A-line mechanical scans, 2D images could be reconstructed. This is called B-mode imaging. Imaged could be formed by doing a linear scan along one direction, or doing a rotating scan. Figure 2-6: Basic principle of the ultrasound pulse-echo and the acquisition of an A-line [16]. There are a few important imaging criteria to evaluate the quality of ultrasound 31 imaging, which are not only applied to mechanical scan image, but also to electronically beam formed image and contract-enhanced image. -- Resolution of the image was defined by the size of point spread function, obtained by imaging a point target embedded in a homogeneous medium. It can also be roughly estimated by calculating the beam diameter given by Equation 2-15. Transducer with high center frequency could produce high resolution. -- Imaging depth: attenuation and absorption coefficient are positively correlated to the wave frequency. The higher frequency used to image, the smaller the imaging depth is reachable. -- Contrast of an imaged is defined by: 𝛶 = z _ 9z } z _ (2-20) where g0 and g m denote the gray levels of the object and background. -- Speckles are the granular appearance in the B-mode ultrasonic image. It is caused by the interference of wavelets scattered by the tissue particles. Strictly speaking it is not a quantitative image quality measurement. Speckles degrades the spatial resolution by obscure small objects. However, speckles do provide information for tissue texture, sometimes used for clinical clinician to make diagnosis. 2.3.2 Beamforming for Array Imaging The main advantage of array over single element transducer is the ability to do dynamic beam focusing for both transmits and receives. During transmitting, the focus, shape and angle of the beam can be controlled electronically by applying appropriate 32 time delay and amplitude modulation to individual element excitation wave; during receiving, the signal from any point within the imaging plane could be obtained by delay, weighting and summing across echo signal from different channels. The process of delay, weight and integration crossing different elements within an array is called beamforming. More-over, electronically controlled beamforming is much faster than mechanical scan, which improves the speed of ultrasound imaging in a great extent. An illustration of the two primary transmit beamforming methods are shown in Figure 2-7, namely linear scan and phase steering. Linear scan forms scan lines by using a sub-aperture of array elements that slides across the face of the array. Each sub-aperture forms a single scan line, and is then translated by one array element to the next position, producing a rectangular image field. On the other side, phased steering generates angular scan lines by using all elements at once to steer the main beam at different angles. Depends on the type of scan, different array transducer design have to be used. For example, array transducer doing phase steering should have grating its lobes at angles greater than ±90°, in order to increase the field of view (FOV) of the image. This requires the pitch of the transducer to be smaller than 0.5λ. In fact, we use to name the transducer linear array or phase array depends on their application. 33 Figure 2-7: Linear scan and phase steering scan. (A) Electronic focusing by applying delay. (B) Linear scan. (C) Electronic steering and focusing by applying delay. (D) Phase steering scan. [21] When the ultrasound wave propagates through the tissue, part of the beam energy is reflected by the inner discontinuities and return to the phased array. To reconstruct the information from any point object, proper delay needs to be applied. The receive delay for the n th element relative to the transducer origin can be expressed as: 𝜏 c = ∆P 5 = 9L 5 = 9L< ) , <L , 9=) L 5 (2-21) 34 The representation of the geometric relation is given in Figure 2-8. Note that r j instead of r is used for the distant from point object to the center of transducer, because for each steering angle, we used to acquired 1000~10000 point. By calculating the relative distance for each transducer element, the point object can be reconstruct by delay-and- sum. Figure 2-8: Receive beamforming on a phase array. [21] 2.4 Imaging System The block diagram for a duplex ultrasound imaging system designed by Texas Instrument Inc. is demonstrated in Figure 2-9. While details about the array beamformer will be discussed in Chapter 5, critical component for single channel digital ultrasounds transmit/receive (T/R) system is circled in the image. A processer is needed for timing control, data communication, signal processing and generating transmit pulse/signal. The pulser/power amplifier then amplifies the driving signal to higher voltage to excite the transducer. A structure called T/R switch controls the switching between transmitting and 35 receiving to/from the transducer. The returned echo signals needs to be amplified and digitized before processed by the processer. For display, echo signal needs to be filtered, and the echo amplitude enveloped needs to be extracted and logarithmzed. Figure 2-9: Block diagram of a digitized ultrasound Duplex image system, with power and spectral Doppler function. For single element transducer mechanical scanning, only the components circled in the red are necessary. [22] 36 Chapter 3 Principle of Super-Harmonics Imaging In this chapter, the linear and nonlinear model of MCA will be introduced. Simulation of bubble behavior under the super-harmonic experiment condition will be provided. At the end, different modality of super-harmonic image setup and their benefit and limitation will be discussed. 3.1 Microbubble Contrast Agent Linear Modeling Gaseous bubbles without shell has very short lifetime in the order of a few second, which cannot be used as contrast agent. However, the behavior of gas bubble provides the basic model for other bubble structure, thus remains interesting to be reviewed. The resonance frequency for a free air bubble (without shell) of radius size a is given by: 𝑓 L = B =i N _ D (3-1) where γ is a heat constant equals 1.4 for air, P 0 was the hydrostatic ambient pressure equals 1.013x10 5 Pa, 𝜌 - is the density of medium. Under the condition that bubble size being much smaller than the wavelength, the bubble wall displacement being much smaller than its radius, and the surrounding fluid being incompressible, the scattering cross-section of an oscillating bubble is given by: 𝜎 7 = >i , ( 2 , , 9B) , < , (3-2) 37 where f is the incident frequency , 𝜒 is a damping constant consisting of damping, thermal conductivity and shear viscosity properties of surrounding medium[23]. The attenuation coefficient of the bubble medium is given by: 𝛼 = c S = (3-3) where n is the bubble concentration in unit of bubbles per unit volume. Encapsulated gas bubbles have longer lifetime, which are used commercially. Two commercial products, Albunex, and Optison, use the plasma protein as the shell material. Albunex contains air and Optison contains a large molecule inert gas, thus yielding a much longer bubble lifetime. The effect of a shell on the resonant behavior of bubbles needs to be taken into consideration, which may increase the resonant frequency and decrease the scattering cross-section [24, 25]. Scattering cross-section vs. frequency function of bubble with different sizes is given in Figure 3-1 (A). Compared the red blood cell (1.1x10 -12 cm 2 =1.1x10 -4 µm 2 ), the scattering cross-section of gas microbubble is 10 7 larger. This dramatically improvement of scattering founds the basement for contrasted- enhanced ultrasound imaging. It is very difficult to produce bubble with uniform size. Knowing the distribution of the bubble diameter and the volume concentration of bubble, the mean scattering cross-section of bubbles can be calculated from: <𝜎 7 >= S i ,y _ c i Pi (3-4) where 𝜎 7 𝑎,𝑓 is the scattering cross-section of a bubble with a radius at frequency 𝑓; N is the total number of bubble; 𝑛 𝑎 𝑑𝑎 is the number of bubbles with a radius between 𝑎 and 𝑎+𝑑𝑎 (Morse and Ingard, 1968). Since scattering cross-section is hard to be 38 measured directly, we can infer by measuring the attenuation, which is dominated by scattering. The experiment and theoretic value of attenuation coefficient of Albunex as a function of frequency is shown in Figure 3-1 (B). Figure 3-1 (A) Calculation of scattering cross-section of a gas bubble, as function of transmit frequency and different radius. (B) Experiment data and theoretical curve of attenuation coefficient as a function of frequency for Albunex. [24] 3.2 Microbubble Contrast Agent Nonlinear Modeling The nonlinear response of a spherical bubble oscillation behavior to a time- varying pressure field can be described by Rayleigh-Plesset equation [26, 27]. de Jong et al [24] , taken the shell into consideration, including the internal friction loss and the restoring force of shell stiffness: 𝜌 - 𝑎𝑎+ 3 2 𝜌 - 𝑎 = =𝑃 z ( 𝑎 G 𝑎 ) +𝑃 −𝑃 G − 2𝛿 𝑎 − 𝑆 7 4𝜋 1 𝑎 G − 1 𝑎 −𝜒 . 𝜔𝜌 - 𝑎𝑎−𝑃 i cos𝜔𝑡 (3-5) where P g is the initial internal pressure of the bubble; P v is the vapor pressure; 𝛿 is the surface tension, P a is the incident pressure, 𝜒 . is the modified damping constant that 39 accounts for the damping caused by shell friction. S shell is the shell elasticity parameter. Solution of the equation gives the time-radius relationship of the bubble. The nonlinear behavior the free-bubble and commercial Albunex bubble is shown in Figure 3-2. Multicycle sinusoid signal at the resonance frequency (both at 2MHz) is used to drive. Due to its larger stiffness, the oscillation of Albunex is weaker compared to the resonant free-bubble. The 1 st and 2 nd order scattering cross-section as a function of transmit frequency is shown in Figure 3-3, in which the 1 st order curve can be calculated from Equation 3-2. Figure 3-2 The change in diameter (left) and in scattering-cross section (right) of free- bubble(up) and Albunex. Acoustic driving frequency of 2MHz at the pressure amplitude of 50kPa is given. [24] 40 Figure 3-3 1 st and 2 nd scattering cross-section as a function of driving frequency, left: free gas bubble; Right: Albunex. [24] Before discussing about the principle of super-harmonics imaging, a summary from the theoretical analysis can be given: - Scattering of MCA is much more stronger than that of tissue and blood; - The resonance frequency of MCA is negatively correlated to its size; - At the resonance frequency the MCA produces the strongest scattering effect; - MCA generates nonlinear harmonic components when excited near the resonance frequency; - The nonlinear behavior of MCA is closely related to the property of shell and core. 3.3 Principle of Super-harmonics Imaging As discussed above, MCAs do not only produce fundamental frequency echoes when excited, as tissue normally did, but also generates sub- or super- harmonic signal due to their nonlinear behavior under acoustic wave [28]. By rejecting the fundamental 41 frequency components, the echo signal from the tissue can be suppressed while the nonlinear harmonic signal from MCAs is preserved, making it possible to distinguish blood vessel structure from other tissue [29]. Super-harmonic imaging, or acoustic angiography, is developed based on this theory and demonstrated to be promising for microvascular structure assessment [11, 30]. It uses high-order (>3 rd order) harmonic signal originated from the MCA to provide high contrast between tissue and blood. Because of the high receiving frequency, typically 30~40MHz, resolution and tissue rejection rate can be improved, compared to other non- linear contrast-enhanced imaging method such as sub-harmonic or lower-order harmonic imaging [31, 32]. Typical MCAs on the market have a median diameter around 1-5µm, with their resonance frequency from 1 to 7MHz. Super-harmonic imaging requires the utilization of dual-element dual-frequency transducer, due to the difficulty of making a single transducer that has strong transmitting efficiency at the resonance frequency while maintaining good sensitivity at the 4th- or higher-order frequency. Such dual-element dual-frequency transducer has been developed in previous studies [10, 30]. 3.3.1 MCA Preparation The micro-bubble used in the current study was provided by Dayton’s research group from University of Northern Carolina, Chapel Hill. The bubble was created using a 9:1 molar ratio of DSPC and DSPE-PEG2000 (Avanti Polar Lipids, Alabaster, AL). These lipid solutions were mixed into 1 mg/mL buffer solutions made of a 0.8:0.15:0.05 ratio of phosphate buffered saline, propylene glycol, and glycerol, respectively. Aliquots of 1.5 mL of were transferred to 3 mL vials, capped, and de-cafluorobutane was ex- 42 changed with the air in the vial headspace using a custom vacuum apparatus. The vials were then stored at 5°C until immediately prior to injection. Prior to injection, the vial was shaken vigorously for 45 seconds using a mixer (Vialmix, Bristol-Myers Squibb Medical Imaging, North Billerica, MA) to produce the microbubble contrast agents. Microbubbles of different sizes were isolated using centrifugation method in the literature[33, 34]. Figure 3-4: Left: Vials of lipid and gas before and after mixing. The white froth is the suspension of microbubbles. Middle: view of microbubble under the microscope. Right: A cartoon shows the structure of the bubble and its lipid shell. [11] 3.3.2 Behavior Simulation The behavior of MCA under oscillating acoustic field was simulated using Matlab (Mathwork, MA) based software Bubblesim, developed by Lars Hoff[35]. The software utilized the Rayleigh-Plesset, de Jong model as presented in Equation 3-5, as well as the linear model introduced in Section 3.1, and solved the bubble radius changing using ODE (ordinary differential equation) packaged in Matlab. If not mentioned, the following parameters were used for simulation: Rayleigh- Plesset, de Jong was used to model the nonlinear behavior, considering shell radiation damping. Medium was selected as water. The transmit frequency was below 10MHz, to 43 be consistent to the super-harmonic application. While in actual experiment the pressure delivered to the ROI could be higher that 1MPa, during the simulation it was only used up to 0.5MPa in order to avoid problem with the ODE solver. 10 cycle cosine tapered sine wave was used to simulate. Variable-order ODE solver based on the numerical differentiation formulas was used. The bubble had a median diameter of 0.8µm. Ø Linear behavior of the MCA denoted by Eqs. 3-1 and 3-2 was simulated. The result was given as the transfer function between the scattered acoustic pressure (Ps) and applied pressure (Pa), shown in Figure 3-5. The relationship between the transfer function and the scattering cross-section was: σ 7 = 4𝜋𝑅 = 1 S 1 3 = 4𝜋𝑅 = ( 7 i ) = (3-6) As predicted, bubble with bigger size had lower resonance frequency. Taking the 4𝜋𝑅 = term into consideration, smaller-bubble had much lower scattering cross- section compared to the bigger ones. 44 Figure 3-5: The transfer function of microbubble with shell of different diameter. Ø Nonlinear behavior of the MCA was simulated. The driving signal and the radius changing of 0.8um diameter bubble is shown in Figure 3-6. The spectrum from 0- 40MHz of the scattered pressure of different bubble diameter are shown in Figure 3-7 [reference to be added]. The amplitude of the graph are in dB (-120dB~ - 180dB), which is relative to the amplitude of the transmit pulse train. For bubble of 0.4 and 0.8µm diameter, the nonlinear response at the harmonic frequencies are within -10dB range compared to the 5MHz baseband response. In the second graph from the left for bubble with 0.8µm diameter, resonance does not only occur at i*5MHz (i=1,2,3…), but also at 0.5i*5MHz (i=2,3,4…), as harmonic 45 distortion. This effect became more obvious for bubble sized >1.6 µm, where harmonics components could not be seen. Figure 3-6: Driving pulse and the bubble radius oscillation. 0.8µm diameter bubble. Ø Population response: given the bubble diameter, using Equation 3-4 and Equation3-6, the mean scattering cross-section could be computed by Equation 3-7. < 𝜎 7 > = 𝜎 7 𝑎,𝑓 ¤ G 𝑛 𝑎 𝑑𝑎 𝑁 = 𝜎 7 𝑎,𝑓,𝑖 c § ∙𝑝 𝑖 =4𝜋 𝑅 § = 𝑃𝑠 § 𝑃𝑎 = c § ∙𝑝 𝑖 =∁ 𝑅 § = ∙𝑝 𝑖 ∙(𝑃𝑠 § c § ) = (3-7) 46 Figure 3-7: The spectrum of scattered pressure of bubble with different diameters. The x- axis is the frequency from 0-40Mhz, y-axis is the relative amplitude to the driving pulse train, from -120 to -180dB. The blue curves are the driving pulse spectra for reference. 47 where p(i) is the percentage of microbubble of specific size R i , C is a constant considering the same incident driving pulse train were applied, Ps i is the scattered pressure calculated from the simulation of bubble with size R i . The distribution of the size bubble was regroup into 15 bins (n=15), with a step of 0.1µm, shown in Figure 3-8, as well as the mean scattering cross-section. Figure 3-8: Left: bubble diameter distribution; Right: The mean scattering cross-section spectrum of the bubble population the magnitude is normalized to dB relative to the 5MHz component. Comparing to the low-order harmonics (≤3 rd order), super-harmonic scattering cross-section near 30MHz is 5-10dB lower but still remaining distinguishable. In addition to that, considering the stronger attenuation to high-frequency wave and an image depth around 1cm, as shown in Figure 2-2, an addition 10dB weaker signal for the high- frequency component is expected. This raises the challenge on both the transducer and system for super-harmonic imaging. 48 Chapter 4 Configurable Dual-Channel System for Super- Harmonic Imaging In this chapter, the design, fabrication, characterization and imaging experiment result for a dual-channel system for super-harmonic ultrasound imaging will be given. A discussion about ways to improving dynamic range of super-harmonic imaging is given at the end of chapter. The hardware and logic design strategy used in the system proposed in chapter is a guidance for array system with larger number of channels. The content of this chapter has been published before in IEEE Trans. Biomed. Eng. on Dec 8 2015, see reference [36]. 4.1 General Design Consideration To obtain super-harmonic contrast enhanced imaging and B-mode imaging, two identical and interchangeable transmit/receive channels were included to work with dual- frequency dual-element IVUS transducer. For super-harmonic imaging, low-frequency transducer was used to transmit the excitation pulse, and high-frequency transducer was used to receive the echo. For B-imaging, only the high-frequency transducer was used for transmit/receive. On the receiving side, we needed high precision, low-distortion and low noise receiving amplifier to improve the dynamics range and suppress the baseline noise, with a target SNR of 50dB. Variable gain was also needed considering the small amplitude of super-harmonic. The designed bandwidth of the receiving channel was from 1MHz to 49 70MHz, which covered the fundamental to the 7 th harmonic. Electrical impedance matching of 50 Ohm was used. On the transmit side, for B-mode imaging, the transmitter should be able to provide broad-band pulse with adjustable center frequency (e.g. 35MHz) for pulse-echo test. For super-harmonic image, in order to induce detectable nonlinear microbubble oscillation, the transmit pulser also needed to generated multicycle burst, with center frequency corresponding to the resonance frequency of major MCAs available on the market, which is 1-7MHz. Considering the variance among transducers transmission efficiency, high voltage output was necessary. The block diagram of the configurable dual-frequency dual-channel imaging system is shown in Figure 4-1 A. The system was controlled by a FPGA. Bipolar pulses or single frequency burst (pulse train) could be generated inside the FPGA, which then passed to the transmit pulser. A digital T/R switch controlled the system to work in either dual-element mode for acoustic angiography or B-mode. After amplified and filtered by the analog receiving circuit, the echo signal was digitized by a 12-bit 200MHz analog-to- digital converter (ADC). Separate AC-to-DC power supplies were used to provide the analog/digital working voltage (1V, 2.5V, 3.3V and 5V) and the pulse generator voltage (up to 200Vpp). To be synchronized with external motor, general purpose I/O pins were provided. The system communicated to the PC through a PCI-express digital I/O card. A summary of the important device parameters are listed in Table 4-1. The system is depicted in Figure 4-1B, based on a small-form-factor 6-layer printed circuit board (16x20 cm, 0.8 lbm). Two SMA connectors were used for transducer connection. 50 Figure 4-1: Up. Block diagram of the configurable dual-frequency super-harmonic imaging system. Different color shows different power domain: blue: digital supply; green: high voltage for pulser; red, analog supply; purple, low-noise analog DC supply for ADCs. Bottom: Photography of the system. 51 Table 4-1 Devices Used in the Proposed System Function Block Device Parameters Manufacture Specification Pulse Generator TC6320 MOSFET Pair a Max Voltage 200Vpp Rise/Falling Time 15ns Amplifers AD8331 Low Noise, Variable Gain Amplifier b Gain -4.5-55dB -6dB Bandwidth 100MHz Analog-to- Digital Converter AD9230-210 b Sampling 210MHz Resolution 12bits Input Range 1.25V Field Programmable Gate Array Virtex-6 XC6VLX75T c Number of IOs 240 Number of Logic Slices 11640 Data Transfer PCIe Digital I/O Card PXIe6537B d Sampling Rate 50MSPS Bus Width 40bits/Sampl e DC-DC converter LT3083, LT3633 e Total Power Consumption 3W,12V a Supertex Inc., Sunnyvale, CA. b Analog Devices, Norwood, MA. c Xilinx, San Jose, CA. d National Instrument, Austin, TX. e Linear Technology, Milpitas, CA. 4.2 Specific System Design 4.2.1 FPGA Logic A Virtex-6 FPGA served as the central controller of the proposed system, working at 200MHz clock rate. It received commands from the PC, controlled the pulse generator waveform and transmitted the data to the PC for display and further processing. The PC communicated to the FPGA through a PCI-express 40-channel digital I/O card, with 50 MHz clock rate, which could be programmed using LabVIEW software (National Instrument, Austin, TX). 52 The digital logic diagram of the FPGA is shown in Figure 4-2(A). To collect one frame of data, an Acq_Start signal was sent to the FPGA from PC through the digital I/O, together with several bits of control signals. Major user-defined control signals are listed in Table 4-2. The mode switch was implemented on the FPGA, which controlled the on-and-off of the analog amplifier, and selected the channel for echo signal receiving. Figure 4-2 (B) shows the timing diagram of image acquisition. At the rising edge of signal Pulse_Trigger, digital pulse/pulse-train were sent out through one pair of differential signal pins on the FPGA, which were then amplified by the pulse generator to get enough excitation voltage. We used single pulse excitation for B- mode imaging and multi-cycle burst for acoustic angiography, which was generated inside FPGA in the form a differential digital wave. The digital pulse was then given to the transmit pulser. After echo signal was received, optional processing such as finite impulse response filtering, envelope detection was implemented on-board. Raw/processed data was stored into a First-In-First-Out buffer. Once full, a FULL bit was generated, acknowledging the digital I/O card for data transferring of a line of data. A LabVIEW graphical user interface (GUI) was developed for control and data acquisition. The software logic diagram and the GUI are shown in Figure 4-3 . 53 Table 4-2 Major User Defined Parameters of the Proposed System Parameter Selection Control Signal(s) Imaging Depth 2048/4096/8192 sample points/line Pre- loaded in FPGA Number of Lines in a Frame N:1-10000 #of_Lines [0-7] Number Pulses Repetition in Each Line 1-2000 #of_Pulse s[0-6] Transmit Pulse Polarity 0° or 180° Control [0] Transmit Pulse Frequency Any Integer Division of the Clock Frequency Control [1-7] Stimulation Mode Pulse Echo Mode/ Dual-element Mode Mode Receiving Channel Chan1/Chan2 Channel Transfer Clock Speed ≤50MHz Pre- loaded in FPGA 54 Figure 4-2: (A) FPGA logic block diagram. Dashed lines show the clock signal. (B) Logic timing of one frame data acquisition. Pulse number 1 and 3 are bipolar pulses for B-mode imaging; Pulse number 2 and 4 are multi-cycle pulse-train for super-harmonics imaging. 55 Figure 4-3: (A) LabVIEW data acquisition software block diagram. Dashed lines indicated the concurrent stages. (B) GUI for raw data acquisition: 2000line/frame, 100Hz PRF, online software filtering and pulse echo mode. 4.2.2 Analog Receiver Design The echo signal from each transducer was passed into an individual analog receiving channel. To protect the receiver circuit, the signal first went through a customed passive limiter, which bypassed the high amplitude transmits pulse to the ground [37]. To provide enough gain for the low-amplitude super-harmonic signal, the echo was then amplified by a low-noise variable-gain amplifier (AD8331), with overall adjustable gain of -4.5 dB to 54dB. The -3 dB bandwidth of the amplifier spanned from 1 to 100 MHz when 50 Ohm input impedance was used. The amplifier also converted the single-ended input to differential output, which eliminated the DC 56 noise. Echo signals were then low-pass filtered by a 3 rd order Butterworth filter, with - 6dB cut-off frequency at 60MHz. The signal was digitized by a 12-bit ADC AD9230 at a sampling rate of 200MHz. After converting, the differential digital signals were routed to the FPGA. 4.2.3 Pulse Generator The transmit pulse generator was able to produce high-voltage bipolar pulses for B-mode imaging, and burst (square wave train) for acoustic angiography imaging. Two MOSFET drivers (ISL55110, Intersil Corp., Milpitas, CA) were used to amplify the digital pulses generated by the FPGA to 20Vpp. The high-speed MOSFET pair TC6320 was used as the voltage amplifier for pulser, which produced bipolar pulses with variable output voltage from 30Vpp to 200Vpp [38]. The amplitude could be adjusted by changing the DC supply voltage. The pulses/burst were passed through two stages of expanders before sending to the transducer, implemented by PMDB7000 high-speed, high-breakdown voltage diode (NXP Semiconductors, Eindhoven, Netherlands). For super-harmonic imaging, multiple-cycle burst need be used to drive the low- frequency transducer in order to induce the MCA nonlinear oscillation. Considering the complexity of designing a high-voltage-gain high-speed linear power amplifier, the same pulser was used for multi-cycle wave generation. Thus the output of the pulser was actually a square wave instead of single frequency sine wave. We will show in the result section that the square wave won’t affect the performance of super-harmonic imaging. 57 4.3 Transducer Design For super-harmonic imaging, it was essential for the excitation wave and the sensing wave to be confocal. In order to make the transducer small enough to fit into the intravascular catheter while keeping the aligned beam of the dual frequency waves, the stacked layer design was selected for this application [10]. The transducer was designed as multiple layers with the high frequency piezoelectric layer in front of the low frequency piezoelectric layer, see Figure 4-4(A). The low frequency excitation wave was generated by the low frequency element (3 mm x 0.6 mm x 0.3 mm), and the wave could penetrate the high frequency element to propagate to the front medium. The high frequency super-harmonic generated by microbubbles was detected by the high frequency reception layer (0.5 mm x 0.6 mm x 0.065 mm), which was in front of the low frequency element. An acoustic filter layer was sandwiched between the two piezoelectric layer, which enhanced the low frequency wave penetration efficiency and reflected the high frequency wave so that no aliasing echo was generated[39]. The whole transducer (4 mm x 0.6 mm x 0.5 mm) was mounted on a 20 Gauge hypodermic needle to mimic the housing for intravascular applications. Eventually, the transducer was coated by parylene with 15 µm thickness, which acted as both shielding and the matching layer of the high frequency element. Common ground was shared between the transmission and reception elements, while the low frequency excitation and the high frequency reception signals were connected individually. A photograph of the transducer is shown in Figure 4-4(B). Detailed fabrication procedure was reported previously and summarized in Table 4-3 [12]. 58 Table 4-3: Transducer Used For Single-element Super-harmonic Imaging Parameter Dual-Frequency Dual-Element IVUS Transducer Low Frequency Element High Frequency Element Center Frequency 6MHz 35MHz Active Material PMN-PT PMT-PT Thickness 300µm 65µm Matching Material Al 2 O 3 /epoxy Parylene Backing Material Silver epoxy N/A Aperture Size 3x0.6mm 0.5x0.6mm a Lithium niobate, Boston Piezo-optics, Bellingham, MA. Figure 4-4: (A) Acoustic stack of the dual-layer transducer. (B) Photo of the transducer front looking. The low frequency element is the 4mm dark color block. The high- frequency element is the light color piece at the center. (C) Experiment setup for vasa vasorum phantom imaging. (D) Microbubble size distribution. 59 4.4 Experiment Setup In vitro IVUS imaging of heathy rabbit coronary arteries was conducted to demonstrate the B-mode image capability. The system was operating in the pulse-echo mode, with the high-frequency element of the dual-frequency IVUS transducer connected to one of the two channels. The pulse generator provided a single cycle of bipolar pulse with 60V pp and 33MHz center frequency. The transducer was attached to a customed rotational motor. For super-harmonic contrast imaging, both high-frequency element and the low- frequency element were connected to the system. Transducer was attached to the same step motor and was placed in a 4mm diameter channel cut through a tissue mimic phantom. The phantom was prepared using a 0.075 g/mL concentration of 275 bloom gelatin with a 0.032 g/mL concentration of graphite scatters. The speed of sound was 1532m/s. The acoustic attenuation coefficient was expressed by the following Equation 2-6, with unit of dB/cm. In 4.5-7.5MHz range, it was measured 𝛼 G =1.399, γ=0.9439 (R2=0.952); in 15-30MHz range, 𝛼 G =0.1228 and γ=1.4172 (R2=0.982).To mimic the size of vasa vasorum, we inserted acoustically transparent tubes with 200 µm diameter through the phantom. The tube was transverse to the transducer imaging plane, as shown in Figure 4-4(C). Detail about the phantom making could be found from the study of Madsen. Polydisperse lipid-shelled microbubbles were formulated as described in the literature. Diluted microbubbles with a concentration of 0.8x10 8 microbubble/mL were pumped through the tube at a mean velocity of 4.4 cm/s while imaging. The microbubble 60 had a median diameter of 0.81 µm, whose resonance frequency was around 6MHz. The microbubble size distribution was shown in Figure 4-4(D). Raw RF data was acquired in the current study. Data was post-processed in MATLAB (Mathwork, Natick, MA) using a 25-40 MHz 6 th order band-pass Butterworth digital filter. Envelop detection was applied for image displaying using a Hilbert transform. 4.5 System Characterization 4.5.1 Analog Receiver Characterization The analog receiver’s impulse responses were test. A pulse generated by Panametrics 5900PR was delivered through the SMA connector to transducer with 40dB attenuation. The pulse waveform and the received signal collected by the proposed system are shown in Figure 4-5. From the time domain waveform, it was evident that the 2 V rms (root-mean-square) baseline noise measured was removed by the analog receiving circuit. The frequency spectrum showed the analog receiver had a -6 dB cut-off frequency at 52 MHz, which was lower than the designed frequency of 60 MHz. The reason for the miss-match between the design and actually frequency was that a few passive components were not inserted at the input stage of the preamplifier in order to match the actually transducer impedance. More specifications were measured. ADC SNR was 55.9dB, which was defined by: 61 𝑆𝑁𝑅 P© =20𝑙𝑜𝑔 BG ( S3®¯° ±3S² ) (4-1) ,where Asignal was the full scale of ADC(1.25V) and Anoise was measured by powering on the system but without signal input. The noise level of the amplifier was tested by measuring the minimum detachable signal, where sin-wave was generated by the function generator and attenuated by series of RF attenuators, which was then sent to the receiving channel of the system. At 35MHz, the minimum detectable signal was 65 µV. The gain at 35MHz was 42.9dB. Figure 4-5: Impulse response: tested by applying a short pulse from a pulse generator Panametrics 5900PR to the transducer connector, and collect the digitized signal from the proposed system. (A) Input pulse voltage (V) collected by the 2 GHz digital oscilloscope (dashed black) and the received signal (scaled to 1, blue) collected by the system. (B) The spectrum of the input pulse (dashed black), of the received signal (blue) and of analog filter simulation (red). 62 4.5.2 Pulser Characterization The output of the transmit pulse generator was measured by a digital oscilloscope (TDS 5052, Tektronix, OR). Figure 4-6(A)(B) shows a pulse centered at 30MHz and its spectrum, and the pulse echo signal collected from the 35MHz element transducer. The - 6dB bandwidth of the pulse covered from 8MHz to 50MHz. Center frequency of the high-frequency element was 35.6MHz, -6dB bandwidth was 47%. For the low-frequency element, the actually center frequency is 5.8 MHz and -6dB bandwidth is 48%. The pulser could generate burst signal (square wave) for super-harmonic imaging. As shown in Figure 4-6(C), the spectrum of a 3-cycle square wave had strong odd harmonics, which were 10dB weaker every 2 octaves. The acoustic output of the 6MHz element driven by the burst was also measured, using a needle hydrophone (HNA-0400, Onda Corp., Sunnyvale, CA) placed axially 3 mm away from the transducer. Shown in Figure 4-6(D), the odd-order-harmonic signal originated from the transmit pulse train was suppressed by the transducer, but there was still a 22MHz high-frequency components in addition to the 6MHz baseband. It was important that the excitation acoustic wave should not have high-frequency leakage, because tissue scattering of the high-frequency would impaired the contrast between bubble and tissue. As shown here, the harmonics of the excitation pulse were filtered by the transducer during transmitting, thus it won’t affect the image quality. However, to suppress the 22MHz component, we needed to apply a filtering with 25MHz low-cut frequency during post image processing. Figure 4-7 shows the relationship between the supply voltage and low-frequency transducer acoustic output. The slope, or the transmit pressure to voltage ratio, was 19.37 kPa/V for a 2-cycle burst case, and 14.46kPa/V for single cycle pulse. Above 70 V pp , 63 both 1-cycle and 2-cycle bursts could produce peak negative pressures over 1 MPa, which was sufficient to produce high frequency super-harmonic of the MCAs for super- harmonic imaging purposes. Figure 4-6: (A) Broad-band pulse and spectrum measured by oscilloscope. (B) 6MHz pulse train and spectrum measured by oscilloscope (C) Pulse echo signal of the 35MHz element collected by the proposed system, shown in normalized amplitude. (D) Acoustic pressure output of the 6MHz element measured by the hydrophone, shown in hydrophone voltage. 64 Figure 4-7: Acoustic pressure collected by the hydrophone, calibrated to pressure. Measured with 6MHz 1-cycle or 2-cycle transmit pulse. 4.6 Imaging Result Figure 4-8: In vitro B-mode image of the rabbit coronary artery obtained by 35MHz transducer element. 65 Figure 4-8 shows the B-mode image of rabbit coronary artery collected using the high-frequency (35MHz) element of the IVUS transducer and the proposed system. 45dB dynamic range was achieved, while the front and back layer of artery were clearly visible in the image. 2000 lines were acquired in this frame. In addition to filtering and enveloped detection, moving average crossing A-lines was applied to this image. Acoustic angiography image was acquired using tissue mimic phantom. The IVUS transducer was placed inside the lumen of phantom. The imaging plane was half way to the bottom of the phantom, where the vasa vasorum mimic tube was about 6mm away from the transducer center. For B-mode image, we injected water into the tube. The result is given in Figure 4-9(A), where the approximate location of the tube is circled, which could hardly be distinguished from the tissue. B-mode imaging with water injected was very similar to Figure 4-9(A), result not shown. Super-harmonic image was acquired using the same setup, but MCAs were injected continuously into the vasa vasorum mimic tube. 3-cycle transmit burst with 120Vpp supply voltage was used. As shown in Figure 4-9(B), the MCAs inside the tube were visualized, while the fundamental frequency containing mostly tissue mimicking phantom backscatter was successfully suppressed. A dynamic range of 30 dB was achieved, which was a 10 dB improvement over previously reported results[10]. Stronger ring-down from transmit wave could be seen near the center of the image. 400 lines were acquired. No moving averaged was applied for the images. 66 Figure 4-10(A) shows a zoom-in view of MCAs inside the tube, re-drawn from Figure 4-9(B). 200µm distance in the axial direction is labeled with a red bar. The size of the tube in the image was around 300um from -6 dB measurements. The vessel mimicking tube was not transparent under super-harmonic mode, probably due to the aging and dehydration of the tube material. There were three major ways to distinguish the MCA signal from the tube signal. First, the super- harmonics from MCAs covered an area that was wider than the tube diameter and had a vague layered structure, while the signal from tube had a clear layered structure representing the tube cross-section. Second, the signal from MCAs was non-uniform across different frames, while that from the tube was uniform. Third, the signal from the tube could be suppressed or eliminated when a small 2-3 °tilting angle was applied to the sample, while the transducer was still able to rotate inside the lumen, but MCA signal was not affected by the tilt angle. The reason why the tube was imaged larger than its actually size was primarily because of the ring-down effect of the transmit transducer, which leaded to aliasing on axial dimension. To demonstrate it, more images were collected using another dual- element dual-frequency transducer, whose low-frequency transducer had 6MHz center frequency and 20% bandwidth, 28% narrowed than the one used for Figure 4-9. Figure 4-10(B) shows the cross-section view collected of the vasa vasorum mimic tube, where the front and back surface of the tube cannot be seen. In Figure 4-10C, another imaging plane was chosen, where the tube runs nearly parallel to the imaging plane, so the longitudinal view of the MCAs carrying tube could be imaged. The sizes of micro-tube in these images were about 500 um. 67 Figure 4-9: Phantom acoustic angiography image. (A) B-mode image, when water is injected into the micro tube. The location of tube is circled with red. (B) Super-harmonic image obtained when MCAs are injected into the micro-tube. 68 Figure 4-10: Expanded view of super-harmonic image. (A) is redraw from Figure 4-9(B). (B) and (C) are collected using a different transducer with narrower bandwidth. (C) is collected where the micro-tube runs parallel to the imaging plane. 200um along the axial direction is indicated by the red bars. 69 4.7 Discussion An integrated, compact, dual-channel imaging system for super-harmonic contrast-enhanced ultrasound was developed. The system supported acquisition of high- frequency B-mode imaging, as well as high-order super-harmonic contrast-enhanced imaging. The measured SNR of the system was 56 dB with adjustable gain. When connected to a dual-element, dual-frequency (6 MHz, 35 MHz) intravascular transducer, the system was able to achieve 45 dB of dynamic range for ex vivo coronary artery imaging. It could also detect contrast agents within a 200 µm vasa vasorum mimicking vessel embedded in an in vitro phantom with a dynamic range of 30 dB, which was a 10 dB improvement over previously reported results. Some of the key design features for this integrated system that directly impacted the quality of the images obtained are discussed below. 4.7.1 Selection of Transmit and Receive Frequency Although it was possible to excite the microbubbles above its resonance frequency and make harmonic images using a single transducer [27, 28], the nonlinear behavior of microbubbles were more pronounced when properly excited near their resonant frequency [50]. As a result, a transmit frequency was selected near the bubble’s resonance frequency, where less pressure might be used for contrast specific imaging, which was ideal since nonlinear signal generation from tissue scales dramatically with transmitted pressure [51]. 70 On the receiving side, higher receiving frequency provided better rejection of the tissue backscatter signal. However, a trade-off had to be made considering the stronger attenuation effect of the high-frequency component and its impact on detectable signal strength. Higher-order harmonic imaging studies have been conducted to optimize which pressures and frequency combinations yielded optimal contrast to tissue ratios, generally noting that lower transmission frequencies and 3rd to 6th-order receiving frequencies produce higher dynamic range and SNR [41]. For the current study, 6 MHz transmitting and 35 MHz receiving was optimal considering both signal strength and backscattering rejection. 4.7.2 Resolution The resolution of B-mode image could be estimated by measuring the -6 dB width of the point spread function, which was 30 µm axially and 110 µm laterally, measured by the proposed system and the high-frequency transducer element (fractional bandwidth 47%), results not shown. For super-harmonic contrast-enhanced imaging, however, because of the flow of MCAs, it was hard to identify a static point target and to define the resolution. As shown in Figure 4-10, the size of the vasa vasorum mimicking tube in the image was larger than its actual size when the transmit transducer with narrower bandwidth was used, which might be explained by ring- down in the excitation waveform and axial aliasing. This implied that the super- harmonic imaging axial resolution was related to both transmit pulse-width and receive element resolution, which was in agreement with the findings in the literature 71 [41]. Thus, in order to improve resolution, transmit and receive transducer with higher bandwidth needed to be used, as well as shorter transmit driving signal. According to the discussion above, imaging resolution could be improved by reducing the number of cycles in the transmit pulses. However, a certain threshold of negative peak pressure had to be reached, in order to excite detectable super- harmonics. Using the current transducer, we could not get any super-harmonic signal using 1-cyle 200 V pp (maximum voltage) pulse. This could be explained by the damping introduced by the transducer. As shown in Figure 4-6(D), when driving by a multi-cycle pulse train, the transmitting negative pressure ramped up and peaked at the 2 nd pulse. If the time constant of the transducer could be reduced, presumably shorter pulse could be used to reach the transmit pressure threshold and thus resolution could be improved. 4.7.3 Dynamic Range The absolute strength of the super-harmonic signal generated from the MCAs nonlinear oscillation was determined by the non-linear oscillation signal amplitude, which was directly related to the acoustic power delivered to the focus of the receive element. As the driving voltage to the transducer went up, the energy delivered to this region increased linearly (see Figure 4-7). Other studies have reported that increasing the transmitted pressure used to excite microbubbles improved dynamic range until a threshold pressure was reached where it was likely that microbubble fragmentation occurred and limited further increases in dynamic range at higher pressures [52]. The 72 loss of echogenicity from microbubbles fundamentally limited the maximum dynamic range obtainable. The second factor that reduced the dynamic range of the system was the higher-order harmonic components that were emitted from the transmitting element. As shown in Figure 4-6(D), there was a 22 MHz component generated by the 6 MHz transducer when driven by a 3-cycle burst. This harmonic component was closed the - 6 dB bandwidth of the receiving element, meaning tissue could reflect higher order harmonics and reduce the dynamic range of the system. This high frequency component from the low-frequency transducer might be originated from the 19.5 MHz harmonic from the driving waveform, as shown from Figure 4-6(B). To get rid of the harmonic, a passive harmonic reduction circuit reported in literature could be used, which could suppress the 3rd-order harmonic by more than 10 dB [53]. Using sinewave generator would minimize the harmonic, but the size and cost-effectiveness would be compromised as mentioned above. A final factor that could be controlled to increase the dynamic range of the system would be the noise level. In the proposed system, a SNR of 56 dB was achieved. However, there was still room to improve the sensitivity and noise tolerance of the onboard amplifier and ADC converter circuits. 4.8 System Upgradation Since only a single element was used for receiving in both imaging modes, it was possible to further reduce the system size by combining the two receiving 73 channels into one using proper multiplexing. Another improvement that could be implemented was real-time imaging function, which would be important for intravascular in vivo imaging, considering the motion of blood vessel in normal heart cycle. By utilizing a data buffering system such as a double data rate type three synchronous dynamic random-access memory (DDR3 -SDRAM), current hardware structure and communication solution was capable to support real-time imaging. The upgraded dual-channel system is shown in shown in Figure 4-11, in which the DDR3 memory was added. There were also a few more triggering in/out signal included. Digital gain control function was provided, enabled by the new pair of DAC blocks. Better power rail design and PCB layout was used. Figure 4-11: Photograph of the second generation of dual-channel system. 74 Chapter 5 Ultrasonic Array System for Super-Harmonics Imaging Array imaging system will be used to acquire super-harmonic imaging in this chapter. Improvement on the system sampling rate is presented. Using the system, single- element-transmit-array-receive mode super-harmonic imaging result will be presented. 5.1 64-channel Ultrasound Beamforming System The system architecture block diagram is given in Figure 5-1, while the photograph of the system is shown in Figure 5-2(A). The transmit subsystem, the receive subsystem and the PC subsystem were built on different PCB board. 64 identical channels were included, in order to work with 64-element phased linear array. Complete detail of the system has been published in [21]. Figure 5-1: Overall high-frequency ultrasonic array system architecture. [21] 64 64-element Array Transducer 64 Digitizer and Receive Beamformer Gigabit Ethernet High Voltage Pulser HVA Pulse Waveform Generator &Transmit Beamformer 64 64 Analog Receiver FPGA Filter DSP FPGA Trigger signals Transmit subsystem LNA T/R Switch 64 ADC 64 Receive subsystem Gigabit NIC Software Application B-mode processing Image visualization Scan convertion Data save PC Subsystem 75 Figure 5-2: (A) Photograph of the developed ultrasonic phased array system. (B) Transmit beamformer. (C) Receive beamformer. [21] 5.1.1 Transmit Beamformer The transmit beamformer (Figure 5-2(B)) was in charge of generating 64-channel delayed output wave. For each scanline only single focus was used. It was developed based on Spartan 3E1600-4fg320 FPGA produced by Xilinx. For each scan line, the delay for every transducer channel was saved as a 10bit binary value. The delay table was generated by Block ROM generator IP Core, with a size of 128 x 640 bits. Finer delay <2ns should be used in order to improve the transmit focus accuracy. Even when 210MHz max clock speed could be reaching (limited by the ADCs), finer timing than 76 single clock period was necessary. Thus, 90°, 180°, 270° phase shifted clocks were used, generated by the Digital Clock Manager IP Core. The core logic diagram in the transmit beamformer is shown in Figure 5-3. For each channel, a pair of positive/negative digital differential signal was generated at the time specify by its delay file. The pair was then passing to an analog pulse generator board, which amplified the signal and delivered it to the transducer. Figure 5-3: Block diagram of the 64-channel transmit beamformer. 5.1.2 Receive Beamformer The receive beamformer (Figure 5-2(C)) was in charge of buffering data and doing delay-and-sum for 16 channels of data (4 boards in all). Dynamic focusing was used for each scan. Based on Virtex 5 VFX70T-1ff1136 FPGA produced by Xilinx. The logic diagram of the receive beamformer was shown on the left of Figure 5-4. The 16 channels of ADC data (12bits/point, 2048 points) was first wrote to Block RAM1 for data buffering; delayed address was generated by the delay computing logic; according to the delay address, data from the 16 channels were read from RAM I, summed and saved in RAM II. The delay computing algorithm was shown in Figure 2-8. For a certain point on 77 the scan line, the distribution time from it to the center of the transducer array (0,0) could be expressed in Equation 2-21. For phase array beamforming algorithm, xn was the transducer position (n=1…16), rj was the distance from the point j=1…2048 to the origin (0,0), and cos α was the scan angle (α have 128 values). Note that τ ´ , the delay file, had a size of 12bit x [2048x128x16], which was too large to be saved in FPGA. By using Tayler expansion, Equation 2-21 could be simplified to: 𝜏 = §∙C∙y S ∙ µ´(8 ¶ ) 5 − §∙C∙y S ∙ 8 ¶ , ·5 , =𝐵𝑆 .. + T¸ ;; · (5-1) as a function of the number of point j. The dimension of delay file was thus reduced to 12bit x [128x16] x2 (BStt and RFtt), which could be saved in FPGA and used to compute the delay for every point. Figure 5-4: Logic Block diagram of the 64-channel receive beamformer. [21] 78 5.1.3 Transducer and System Calibration The array system was calibrated with a 30MHz 64 phase array developed in our lab [20], the specification of a 30MHz phase array Figure 5-5(A), with specification listed in Table 5-1 Figure 5-5: (A) 30MHz phase array. (B) US image of a wire phantom by a 30 MHz PMN-PT kerfless phased array. [21] Table 5-1: Specifications of the 30 MHz prototype linear phased array [21] Specification Values Center Frequency (MHz) 26.3 Bandwidth 58 Number of Elements 64 F number 2.6 Element width in azimuth direction (µm) 24 Kerf width in azimuth direction (µm) 3 Pitch in azimuth direction (µm) .0 Element length in elevation direction (mm) 2 Lens with elevation focus at (mm) 5.5 79 Five 20-µm-diameter tungsten wire targets (California Fine Wire Company, Grover Beach, CA) were used for imaging. The wires were arranged diagonally with the equal distance in the axial (1.5 mm) and lateral (0.65 mm) directions. The image was acquired as the transducer scanned across the wire phantom placed in degassed water. An image of the five-wire phantom is shown in Figure 5-5(B) using a linear gray scale and 50-dB dynamic range. The scan angle was set at ± 30 o . No apodization or thresholding was implemented during reconstruction. The measured full-width half-maximum (FWHM) resolutions were 467 µm and 118 µm for the lateral and axial directions, respectively. 5.2 Sampling Rate Upgradation The SNR of the array system was relatively lower (42dB). One of the reason was the relatively low sampling frequency used. In the current application, 30MHz was the center frequency of echo signal. But for a transducer with 50% bandwidth, which was typical for many high-frequency transducers, 45MHz need to be covered. Usually in ultrasound, we used 2 x Nyquist rate, which was 4x the center frequency of the transducer. This required the ADC to work at nearly 200MHz, as the system demonstrated in Chapter 4, which had 56dB SNR Identical ADC was used in the array system and the dual-channel system, with a max sampling rate reached 210MHz. But the system was only able to sample at 140MHz. The reason was that the ADCs for the receiving beamformer design were synchronized to the FPGA input buffer using a global clock. Thus, the max speed of the sampling was limited by the max throughput of the FPGA to run the beamforming algorithm. Moreover, 80 receive beamformer was synchronized to the transmit beamformer as well. To update the sampling rate to ADC’s full capability, both the transmit and receive algorithm throughput needed to be updated as a whole. - Transmit Beamformer Detail of the delay generation logic is shown in Figure 5-6. For each channel, the higher 8 bits from the delay number were used as the coarse delay number, the lower 2 bits were used as the fine delay number. A counter synchronized to clock 0° (no degree shift) was used to count for coarse delay amount. Once done, a registered done_00 was generated, clocked by clock 0°. Done_01, done_10, done_11 were also registered from done_00, synchronized to clock 90°, clock 180° and clock 270°, respectively. Four pairs of positive/negative digital signal were generated, with 90° shift to each other. Depends on the value of the fine delay, one of the four output pairs was selected through a combinational mux. Setup violation occurred at 200MHz, because when the done_00 signal was ready after the clk 0° rising edge, it was less than ¼ clock period to the rising edge of clk 90°. It is observed that majority of setup violation happened to clk 90°, less to clk 180°, and least to clk 270°. 81 Figure 5-6: Transmit delay generation algorithm. Setup violation occurs at the Ө sign. 82 Figure 5-7: Modified transmit delay generation algorithm. Figure 5-8: Delay computation algorithm. 83 Avoiding inter-clock domain communication by make four individual copies of clock domain could solve the setup violation problem. However, the parallelism had to begin from the very beginning of the sequential logic, where the delay number was read from the block ROM. Shown in Figure 5-7, instead saving 4 copies of data, 2x dual port ROM was used to reduce the redundancy. The coarse and fine delay number, the done signals and the outputs generation of the four clock domains were then processed in parallel, until the very end. By using this method, no cross phase setup violation happens. The speed of the clock was successfully improved from 6.03ns to 5.05ns (160MHz à 198MHz). - Receive Beamformer As describe in previous section, the delay number of the receive beamformer needed to be computed in real time by doing one division and two adding. The implementation of the algorithm is shown in Figure 5-8, making use of the math IP core provided by Xilinx. The throughput of the divider and adder were both setup to 1 clk. It appeared that fanout problems occurred at the input/output of this delay computation logic, circled in Figure 5-8, which leaded to setup violation. Duplicating the registers reduced the fan-out problem, which solved the problem locally but caused other violations globally. The finally solution was to be adding a register between the output of the delay computing logic and the address input of the RAM I, shown in Figure 5-4, which was originally registered. All the setup violations were solved and the beamformer could now run at 4.76ns (210MHz) compared to before at 6.03ns (160MHz). Another benefit of improving the clock rate was that the AD9230 ADC worked more stable at higher frequency range and produced less noise, which improved the overall SNR of the image. 84 5.3 Single-Transmit-Array-Receive (STAR) Mode Comparable to dual-element mechanical scan approach, dynamic transmit and receive focusing could be the optimal solution to provide a uniform point-by-point focusing for super-harmonic imaging. However, there was no existing dual-frequency array successfully developed which had the wide frequency span required for super- harmonic imaging. Wang et al. [40] build a small factor size lateral mode array, which had a 32-element 30MHz linear array on top of a 2.25MHz transducer with 8 sub-element that connected in parallel. The low frequency element was not able to do dynamic focus and was only able to provide unfocused beam to excite the microbubble. The array was able to acquire super-harmonic data with 12dB contrast using standard delay-and-sum beamforming, lower than the result of dual-element imaging. Consider the complexity of make dual-frequency array, in the current study, we used single-transmit-array-receive approach, which we named STAR mode super- harmonic imaging, as illustrated in Figure 5-9. In STAR mode, a signal element low- frequency transducer (2-6MHz) was used to deliver the high acoustic pressure required for super-harmonic imaging, while a high frequency array transducer (around 30MHz) was used to receive the high-frequency harmonics from the MCAs. The receiving array passively received the super-harmonic echo and did receiving beamforming, either by phase steering or linear scan. 85 Figure 5-9: (A) A cartoon shows single-transmit-array-receive (STAR) super-harmonic imaging. The transmit beam approximation is shown as red dash line. During the array acquisition, the start of each scan line was synchronized with the transmit transducer sending one excitation pulse. The transmit transducer did not move during one frame acquisition. Compared to dual-element mode, STAR mode super- harmonic was able to image much faster because of electronically scan. However, since the transmit transducer was only able to excite the MCAs close to its focal point, only MCAs within the FOV of both transmit and receive transducer could be imaged. After the receiver finishing one scan, the transmit transducer could be moved for another image, to enlarge the region that MCAs being activated. Thus, by moving the transmit transducer, super-harmonic from every point within the FOV of receiving array could be acquired, which made the operation much easier in a (pre)-clinical environment. Transmit transducer with large F# (>1.5) and longer depth-of-focus (Equation 2-16) should to be used, in order to excite a larger area that potential diffused by MCAs. 86 5.4 Phase Array Imaging Result 5.4.1 Experiment Setup We used the 30MHz phased array introduced in Figure 5-5 and Table 5-1, and the 64-channel array system for STAR mode super-harmonic imaging. A 3.5 MHz focused transducer with 6.25mm diameter and F# of 2 was used to deliver the excitation wave. The beam-profile of the transmit transducer is shown Figure 5-10(A). The -6dB depth-of-focus in the axial direction was 7mm, which was able to cover the array FOV. In Figure 5-10(B)(C), the setup of the STAR super-harmonic experiment is shown. The transmit transducer was put on the side the phase, with its beam almost perpendicular to the phase array, in order to minimize the reflection from the vessel mimicking tube. The tube was placed 5mm to 8mm in front of the array. 0.8x10 8 microbubble/mL of MCAs were pumped through the tube at a mean velocity of 4.4 cm/s. The same concentration and speed was used in the rest part of this chapter. The transducers and the phantom were placed in DI water. During the experiment, the transmit excitation wave was produced by function generator and power-amplifier. The excitation pulse/burst could also be programmed into the array system. 2 cycles of sinusoid wave with 70V pp amplitude was used. The transmit signal was synchronized with the array scanline timing, where 100 line with 60° steering angle were acquired. The single element transducer position was moved manually using a 3D micromanipulator, while the receiving array doing B-mode imaging in real-time. When detectable super-harmonic signal was acquired, the B- 87 mode excitation pulse for the array would be turned off so only the low frequency transducer was excited. Raw data was acquired and post processed. Figure 5-10: (A) Simulation of the 3.5MHz single element transducer beam-profile used for phase array super-harmonic imaging. (B) Setup of the experiment, where the single element transduce was put on the side of the phase array. The microbubble filled tube is covered by the focal point of the transmit transducer. (C) Photograph of the phase array system based STAR mode super harmonic imaging. The zoom-in view of the circled region is show on the left. 88 Figure 5-11: (A) B-mode image of the vessel mimicking tube acquired by the phase array while the 3.5MHz single element transducer transmitting. (B) Super-harmonic signal, where excitation signal of the array was turned off. 25dB dynamic range, using two-way distance calculation. (C) Super-harmonic image acquired at 6mm, one-way distance. (D) Super-harmonic image acquired at 8mm, one-way distance. 20dB dynamic range. 89 5.4.2 Imaging Result Figure 5-11 shows the imaging result of the STAR mode super-harmonic imaging using the phase array setup. In (A), dynamic transmit beamforming at 5mm was used for the phase array to acquire B-mode imaging, while the single element 3.5MHz transducer was sending the excitation wave. Both the pulse echo signal from the vessel mimicking tube and the super-harmonic signal from the MCAs was received by the array. In Figure 5-11(B), where the pulse echo transmit signal was turned off, only the super-harmonic signal was left. However, due to the two-way distance calculation, the position of the super-harmonic harmonic signal was misinterpreted, thus appeared to be deeper than the actual position of the tube. Instead of using Equation (2-19), Equation (5-2) should be used for one-way traveling distance calculation: 𝑑 =𝑐∗(𝑡 .Lic7»§. +𝑡 L5§ ) (5-2) in which t transmit is the time of traveling from the transmit transducer to the imaging point, t receive is the time of traveling from the imaging point to the receive transducer element. Figure 5-11(C)(D) shows the images collected when the tube was placed at 6mm and 8mm depth, while one-way traveling distance was used. The same B-mode receiving beamforming method was applied for these images, where each scan line was delayed and summed based on the steering of the transmit beam. In fact, during the actual super- harmonic signal, the transmit transducer did not move when the scanlines were collected. For each scan angle, the transmit condition was the same. Thus, we could see a widespread super-harmonic signal band due to the lack of dynamic transmit focusing. 90 5.5 Linear Array Imaging Result 5.5.1 Experiment Setup A real-time linear array imaging system produced previously in the lab and a 256-element 30MHz linear array transducer was used for STAR mode super-harmonic imaging as well. [41, 42]. The imaging system had similar hardware architecture as the phase array system introduced earlier, but had more number of channels and included electronical multiplexers handling the switching between the transducer elements. A 5.8MHz focused transducer with 9mm diameter and an F# of 1.5 was used for super-harmonic experiment. The beam-profile of the transmit transducer is shown Figure 5-12(A). The -6dB depth-of-focus in the axial direction was 6.5mm. In Figure 5-12(B), the setup of the STAR super-harmonic experiment is shown, similar to the setup for the phase array. The transducers and the phantom were placed in DI water. 2 cycle of sinusoid wave of 22-85 V pp amplitude was used to drive the transmit transducer. The transmit signal was synchronized with the array scanline timing, where 192 line with 50 µm step size was used. 64 element was used as sub-aperture to acquire each scan line. 5.5.2 Imaging Result The super-harmonic imaging result using the linear array is shown in Figure 5-13, where different transmit voltage was used to drive the 5.8 MHz excitation transducer. Delay-and-sum beamforming for linear array was used for the image formation, as 91 illustrated in Figure 2-7 (B) and the reference [41]. It was obvious from Figure 5-13 (A) to (E), that the increase from transmit power increased the amplitude of the super- harmonic generated from the MCAs, thus enhanced the SNR of the image. By matching the transmit voltage to the acoustic pressure relationship shown in Figure 5-13 (F), we found that detectable super-harmonic signal could be generated at the focus transducer focal point at a pressure level of 0.4MPa. It was 50% lower than the threshold suggested in previous literature, where unfocused single element transducer was used for both transmit and receive [43]. The pseudo-enhancement effect was also enhanced due to the increase transmit power, which was the shadowed region behind the center of the bubble image between 5mm to 5.5mm. Figure 5-12: (A) Simulation of the 5.8MHz single element transducer beam-profile used for linear array super-harmonic imaging. (B) Setup of the experiment, where the single element transduce was put on the side of the linear array. The microbubble filled tube is covered by the focal point of the transmit transducer. 92 Figure 5-13: Super-harmonic image collected using the linear array. (A)- (E): transmit voltage 22, 38, 54, 70, 85 V pp respectively. 25dB dynamic range. (F) pressure output vs. transmit voltage for the 5.8 MHz single element transducer. Figure 5-14: Mechanism for pseudo-enhancement. 93 The pseudo-enhancement was an effect induced by the non-linear property of the MCAs, which was previously reported in contrast-enhanced ultrasound literature [44]. The mechanism of the pseudo-enhancement was explained as cartoon in Figure 5-14. When the low frequency wave induced the MCAs non-linear oscillation, the super- harmonics signal was propagated in all direction into the surrounding medium. The tissue and tissue mimicking tube wall propagated this high-frequency signal linearly due to their weak nonlinearly (B/A factor, see Equation (2-9)), while the MACs propagated the wave in a nonlinear fashion. Thus, when the signal from the back-wall of the MACs detected by the receiver in the front, the position appears to be elongated. Pseudo-enhancement could be reduced by lowering the concentration of MCAs. We will show in Chapter 6 that it could also be reduced by beamforming method. Figure 5-15: (A) Axial resolution of super-harmonic image at different transmit voltages. (B) Lateral resolution of super-harmonic image at different transmit voltages. Due to the ‘blinking’ natural of the super-harmonic signal, we used 20 A-line averaging to calculate the axial resolution, and 0.4mm depth averaging to calculating the lateral resolution. The averaged axial and lateral profiles at different transmit voltages is shown in Figure 5-15. The measurement results are listed in Table 5-2. As we can found, 94 the axial resolution of the images were comparable to the size of the tissue mimicking tube, while the lateral resolution was much worse than the axial resolution. The reason of the poor lateral resolution was due to the lack of transmit dynamic focusing. We will discuss the beamforming method to improve the lateral resolution in Chapter 6. Another finding was that the resolution of the imaging enhanced as the transmit voltages increased. Table 5-2: Measurement of axial and lateral resolution of linear array STAR mode super- harmonic imaging at different transmit voltages. Transmit Voltage 22V 38V 54V 70V 85V -6dB Axial (mm) 0.215 0.232 0.132 0.183 0.168 -6dB Lateral (mm) 1.221 1.329 0.670 0.910 0.783 The real-time imaging is made possible because of the using the of the array system, with a highest frame rate of 30 frame/sec. Figure 5-16 shows eight consecutive super-harmonic images before and after the switching-off of the injection pump, with an interval of 0.5s (15 frames). The pump was off coincident to Figure 5-16(B), labeled with red. It could be found that after 2s, the super-harmonic signal turned very weak to be detected, makes the SNR really low. This effect was due to the instability of microbubble under the acoustic field, which were partially broken due to the high transmit power, showing that consistent flow of MCAs was required for super-harmonic imaging. 95 Figure 5-16: Super-harmonic image acquired before and after the microbubble stop pumping through the micro-tube. (A)- (H) consecutive image collected with 0.5s interval. The pump was turned off at the time of (B). 25dB dynamic range. 96 5.6 Summary and Discussion In this chapter, we presented using STAR mode and array transducer to collect super-harmonic image, which is the very first attempt for the method. We used 2.5MHz F#=2 transducer and 5.8MHz F#=1.5 transducer for transmit, and 30MHz phase array and linear array as receiver, both successfully collected super-harmonic image. Compared to single element approach, using the array made real-time super-harmonic imaging possible. By using the single focused transducer for excitation, lower power (0.4MPa) was needed to induce super-harmonic, and better SNR (25dB) was acquired. The lateral resolution of the imaging was not as good as the axial resolution, which was due to the lack of dynamic transmit focusing. Real-time super-harmonic image was collected for the first time, showing the dynamic behavior of microbubble through the vessel. The beamforming method used in this chapter was still the two-way delay-and-sum method for B-mode, which needed to be modified. A method to suppress the pseudo- enhancement effect is also need. 97 Chapter 6 Synthetic Aperture Beamforming for Super Super- harmonic Imaging In this chapter, an introduction of virtual point synthetic aperture (VPSA) beamforming will be given first. Then demonstration of its application to single- element bio-microscopy and to IVUS imaging will be given. The method will also be used for super-harmonic imaging, both single element based and array based STAR mode imaging, in order to improve lateral resolution, to improve SNR and to take care of the pseudo-enhancement, 6.1 Introduction to Virtual Point Synthetic Aperture Beamforming SA beamforming was a method that used low-resolution image collected by small aperture to construct an image of finer spatial resolution, emulating the performance of larger aperture. The method was first used for radar antenna, later being used in ultrasound imaging, including application in non-medical and medical area [45, 46]. The method has been successfully implemented on both single-element- based ultrasound, as well as array-based electrical scan imaging [47-50]. 6.1.1 VPSA for Single Element Imaging For mechanical scan imaging, SA utilized the overlapping of acoustic field from scans adjacent to each other, and combined echo from different emissions to rebuild high-resolution image. A common technique used was called virtual point 98 synthetic aperture (VPSA), where the focal point of the transducer was considered a virtual source [51]. In traditional B-mode imaging, where a single emission was used to acquire the information along the whole axial distance, the underlining assumption was to view the acoustic field of the transducer as a narrow and uniform ‘pencil’ shape. In fact, the narrowest beamwidth was only achieved at the focal distance, while in the off-focus resolution was inversely proportional to the distance from the focal point because of the beam divergence. For highly focused transducer, the smaller the F#, the stronger the divergence was. The concept of conventional Bmode imaging is illustrated in Figure 6-1 (A), where the shape of the transducer and beam profile is circled in dashed line. For the VPSA method, as shown in Figure 6-1 (B), the shape of the beamprofile was considered. A virtual source/detector was assumed to coincident with the position the transducer focus, which produced spherical wave in all direction [52]. The beam narrowing and beam spreading before and after the focal point were simplified to two cone shapes, where the spreading angle was determined by the aperture and F# of the transducer, and the focal distance was determined by using Equation (2-14). For an imaging point in the overlapping acoustic field, as the green point in Figure 6-1 (B), synthetic aperture beamforming was done by delaying and summing the echo signal from adjacent emissions, whose beamprofile covering the imaging point, seeing Equation (6-1). The time-of-flight was determined by the two-way propagation distance from the virtual source to the imaging point, see Equation (6-2) [53]. 𝑆 ¼½ = 𝑅𝐹(𝑖,𝑡−∆𝑡 § ) §¿B (6-1) 99 ∆𝑡 § =2 𝑠𝑔𝑛(𝑧−𝑧 y ) [𝑎𝑏𝑠 𝑧−𝑧 y −𝑟] 𝑐 (6-2) where 𝑧 was the depth of the imaging point , 𝑧 y was the transducer’s focal depth, 𝑟 was the distance from the focal point to the imaging point. N was the number of overlapping scans included, determined by the beamprofile cones. Note the delay should be negative if the synthetic point was before the focal point. The assumption of virtual source/detector was based on the fact that the beam from the transducer surface added coherently at the focus, so the delay was calculated relative to the virtual point instead of the transducer surface [51]. Theoretical deduction of the method could be found in literature [54]. The VPSA method can improve the image SNR as well as the off-focus resolution. One artifact created by the VPSA beamforming was the weakening of image around the focal depth, because the limited number of scan lined that contributed to the synthetic point near the focus, as shown by the numbers in Figure 6-1 (B). Two kinds of weighting function could be used, one was the linear inverse weighting; the other was coherent factor weighting (CFW) [55]: 𝑊𝑡 §ciL =𝑎 1/𝑁 (6-3) 𝑊𝑡 ĸ = 𝑅𝐹(𝑖,𝑡−∆𝑡 𝑖 ) N−1 i=0 2 𝑁 𝑅𝐹|𝑖,𝑡−∆𝑡 𝑖 | 2 𝑁 𝑖=1 (6-4) where a was constant factor of the linear weighing. In case of perfect coherence, as shown in Figure 6-2, 𝑊𝑡 ĸ =1; in case of zero-mean random noise, 𝑊𝑡 ĸ =0. Thus the CFW method had the ability to improve SNR when integrated with the SA method. Both of the weighting methods were used in the current study. 100 Figure 6-1: (A) Concept of traditional mechanical scan B-mode imaging, where the acoustic field of the transducer is showing in dashed line. The red dot denote the transmit focus. The green dots denote the imaging point. (B) Concept of virtual point SA imaging. The number in the image shows the number of overlapping acoustic field. Figure 6-2: Illustration of coherent factor weighting. Left: cases where the delayed synthetic beams are perfectly coherent; Right: case where the synthetic beam are not well coherent. The vertical axis is the depth, and the horizontal axis is the synthesized aperture direction. [56] 6.1.2 VPSA for Array Imaging SA beamforming was also widely used for array-based imaging in order to improve the resolution across all imaging depth. A very general approach was to use single or multiple elements sub-aperture to emulate the defocused spherical wave of 101 signal element pattern, and to use whole or sub-aperture for receiving[48]. The echo signals from the full scan was used to synthesize dynamic transmit and receive focusing, to achieve optimal (on-focus) resolution at every image depth. Compared to traditional dynamic focusing transmit method, using defocused beam and SA beamforming achieved the full resolution at all depth without sacrificing the imaging speed [45]. SA beamforming also made it possible to develop a smaller scale electronic system used to drive array by multiplexing between less number of channels [48, 57]. This was critical for the cost control and operation feasibility for many ultrasound applications, including the super-harmonic imaging. Figure 6-3: Virtual source concept for linear array beamforming. (A) Traditional transmit focusing; (B) spherical defocusing wave transmit used for SA. [58] VPSA method had also been implemented for array-based imaging [52, 58, 59]. In fact, the traditional transmit focusing and the defocusing SA beamforming could both both interpreted using the same virtual source concept [58]. As shown in Figure 6-3 (A), in traditional linear array imaging, the transmit focus of the sub-aperture 102 could be seen as a virtual source. This was similar to the case for single element transducer shown in Figure 6-1 (B). For defocusing transmit mentioned earlier, the virtual source could be seen as behind the transducer element, as in Figure 6-3 (B). By calculating the propagation delay relative to the virtual source rather than to the transducer surface, better resolution could be obtained at all image depths. For implementation, the VPSA method could be accomplished by adding a synthetic beamforming stage on top of the traditional single-focus dynamic focusing beamformer [59]. The two-stage beamforming method was also referred as synthetic aperture sequential beamforming. The VPSA beamforming downscale the system complexity with improved image resolution or SNR. Hardware system has been built and successfully implemented for many applications [60]. For current study, both stages were performed on the Matlab as post-processing. 6.2 Application of VPSA in Bmode Imaging 6.2.1 Application of VPSA to Ultrasound Bio-microscopy One of the usages of high-frequency transducer was to exam the micro-structure of small tissue, which sometimes called ultrasound bio-microscopy. High-frequency highly-focus transducer was usually used in order to improve the resolution for bio- microscopy. Strong divergence of the beam leaded to a narrow depth-of-view, where VPSA method could be used. 103 We used a highly-focused transducer developed previously in our lab, which had a center frequency of 64MHz, F# =0.6 and 3mm diameter, focusing at 1.8mm. The imaging target was a 0.8mm diameter rat aorta tissue. The beam-profile simulation result of the transducer is shown in Figure 6-4 (A), while the imaging result is given Figure 6-4 (B). The beam converged and diverged strongly before and after the focal point, which caused the blurry in the image at the front and back end of the aorta tissue. System noise could be seen near the transducer surface before 1.3 mm depth. Figure 6-4: (A) Simulation of the 64MHz transducer beam-profile. (B) B-mode cross- section image of the rat aorta tissue. After VPSA beamforming was applied, both the linear and coherent weight method was used, the detail of which has been given in Equation (6-3) and (6-4). The weighting function mapping was shown in Figure 6-5. Linear weighting was primarily an inversion of the number of scanlines used to synthesize an image point, which was a regional processing method; while the CFW calculated the coherence between signals used to synthesize each point, giving a point-by-point weighting function. The effect of 104 both weighting function and VPSA method are shown in Figure 6-6 (A) and (B), in which VPSA method successfully enhanced the image quality by reducing the blurring at the front and back end of vessel. The linear weight got rid of the system noise near the transducer surface, but did not eliminate the blurry around the focus; the CFW, instead, got rid of the random noise around the target, but preserved the system noise near the transducer surface. Moreover, CFW also sharpened the speckles of tissue, give a better contrast of the image. Both weighting method had its own benefits and should be selected accordingly under different image context. For comparison, Figure 6-6 (C) shows the compounding image of six pre-beamformer images acquired at six different depths. The method was called B/D scan, as introduced in [54]. It had optimal lateral resolution across the whole depth and could be used as standard image. We found that the VPSA beamformed images captured the shape of the target better than the pre-beamforming image. Figure 6-5: (A) Linear weighting function. (B) CFW. 105 Figure 6-6: VPSA result: (A) VPSA with linear weighting. (B) VPSA with CFW. (C) compounding image using pre-beamforming images acquired at different depth. 50dB dynamic range. 106 6.2.2 Application of VPSA to IVUS Imaging We also used VPSA beamforming for the IVUS imaging enhancement, in which a 40MHz small factor size IVUS transducer was used [61]. Compared to the examples showing in previous section, where the beamforming was performed in Cartesian coordinate, IVUS beamforming was done in polarity coordinate and changes needed to be made considering the angular sweeping of beamprofile across different scans. Another major difference for the beamforming was that the IVUS transducer was unfocused instead of focused. Thus, when we built the beamprofile, a different shape rather than the cones was needed. The result of the IVUS imaging beamforming can be found in Figure 6-7, where pre-beamforming image and post beamforming ones are given. It was obvious, that the VPSA method improved the image SNR, improved the lateral resolution, as well as enhanced the contrast of IVUS image. 107 Figure 6-7: (A) Pre-beamforming IVUS image of rat aorta. (B) VPSA beamformed image of rat aorta. (C) Pre-beamforming IVUS image of size 20µm wire targets at five different depths.(D) VPSA beamformed image of wire targets. (E) Zoom-in view of (C). (F) Zoom-in view of (D). 108 6.3 Application of VPSA for Super-Harmonic Imaging VPSA beamforming mentioned above was applied to super-harmonic imaging for the first time. Compared to tissue B-mode imaging, there were several differences of beamforming for super-harmonic imaging: I. Super-harmonic signal from microbubble was a non-static continuous flow signal, while SA method usually applied to static or slow motion tissue. II. Because of the flowing natural of MCAs, the super-harmonic signal from the MCA-filled tube/vessel was in-homogeneous, while in B-mode image the distribution of tissue scatters were homogeneous. III. Super-harmonic imaging was a dual-transducer application. The miss-match/miss- alignment between the transmit and the receive transducers’ focal points need to be considered during the synthesizing. As facts I and II mentioned, coherence of the echo signals from adjacent emissions was weak compared to B-mode imaging. However, studies in literature has shown that partial correlations still existed in contrast imaging and beamforming method based on the phase coherence could be used to improve the image quality[62, 63]. Moreover, compared to the size of a single microbubble, the size of micro-vascular was usually around 100µm to millimeter range, which was closer to the resolution of the receiving transducer. Thus, when the concentration of MCAs was high, it was reasonable to assume the micro-vessel itself as a homogeneous scatter. Nevertheless, it was expected that the partial coherence signal from the MCAs would reduce the performance of synthetic aperture beamforming. 109 The mis-match and mis-alignment of transmitting and receiving will be discussed in the examples below. Table 6-1: Transducer Used for Super-harmonic Beamforming Study Parameter Dual-Frequency Dual-Element Transcutaneous Transducer Low Frequency Element High Frequency Element Center Frequency 5.5MHz 30MHz Active Material PZT 1-3 Composite Lithium Niobatea Matching Material Al 2 O 3 /epoxy+Parylene Al 2 O 3 /epoxy+Parylene Aperture Size 9mm 5mm -6dB Bandwidth 50.1% 61% Focus 13.1mm a Boston Piezo-optics, Bellingham, MA. 6.3.1 VPSA for Dual-Element Super-Harmonic Imaging We used a transcutaneous dual-element dual-frequency transducer developed in our lab to study the effect the VPSA beamforming on super-harmonic imaging. Table 6-1 gives the parameter for the dual-frequency transducer. The two elements were con- focused at 13.1 mm. The photograph of the transducer is shown in Figure 6-8 (A). Note that for the high-frequency receiver element, the F# was larger than 2. We first acquired B-mode image using the high-frequency element and test the performance of VPSA on it. Tungsten wires phantom with 20 µm diameters was used as target. As shown in Figure 6-8 (B), the wire at 10.5mm and 14.5mm appeared to be blurrier that the ones near the 13.1mm focus. Figure 6-8 (C) shows the VPSA + CFW processed image. Resolution of all five wires has been improved. The ratio the improvement is given in Table 6-2, from 110 which we found that the VPSA method improved the resolution the most at the off-focus points as the theory predicted. Figure 6-8: (A) Photograph of the dual-frequency transducer. (B) B-mode imaging of wire phantom using the 30MHz element. (C) VPSA+CFW beamforming result. 40dB dynamic range. 111 Table 6-2: Resolution Improvement B-mode VPSA+CFW -6dB Lateral Resolution (µm) 10.5mm 11.5mm 12.5mm 13.5mm 14.5mm Raw B-mode 540 145 151 153 169 VPSA+CFW 70 85 143 103 93 Reduction Ratio (0: no reduction 1: reduce to 0) 97% 41% 5% 33% 45% Figure 6-9 gives the signal spectrum collecting using the dual-element transcutaneous transducer, where the 5.5MHz element was transmitting and 30MHz was receiving. The strength of higher-order harmonic signals was much stronger when the tube was filled with MCAs, compared to the cases when the tube was filled with air and water. The super-harmonic signal got bigger when the transmit voltage increased. Figure 6-9: Signal spectrum acquired using 5.5MHz transmitting 30MHz receiving at different transmit voltages. (A) air tube. (B) water tube. (C) MCA-filled tube. 112 The experimental results are shown in Figure 6-10, where the microtube was embedded in the tissue mimicking phantom and 1x10 8 #/ml MCAs was used. The MCAs was pumped at a speed of 3ml/hr. During the beamforming, only the receiving transducer aperture was used to define the divergence of the beam, based on the fact that the receiving high-frequency transducer had a smaller aperture and focus than the transmit transducer. Compared to the B-mode, where wall of the tube was seen, the super-harmonic image was able to show the MCAs inside the tube, giving a better contrast image without tissue backscattering. After the VPSA+CFW, there was roughly a 10dB enhancement on the SNR of the image, similar to the case of B-mode imaging shown in previous sections. The pseudo-enhancement effect located at 12.8mm was eliminated by the beamforming. Pseudo-enhancement was caused by the non-linear propagation of MCAs. It was suppressed by the beamforming method due to its weakly coherent nature. The improvement on the lateral resolution was not significant. 113 Figure 6-10: (A) 30MHz B-mode imaging of the microtube in the tissue phantom. (B) Pre-beamformed super-harmonic image. (C) VPSA+CFW beamformed image. (D) Longitudinal view pre-beamformed super-harmonic image. (E) Longitudinal view VPSA+CFW beamformed image. 40dB dynamic range. 114 6.3.2 VPSA for STAR Mode Super-Harmonic Imaging VPSA beamforming was also applied to the STAR mode super-harmonic imaging. Compared to the single-element approach, STAR mode used different transmit and receive transducer, which should be counted during the synthesis of the aperture. In the first step, traditional B-mode beamforming was used. In the second synthesis step, the delay lines was calculated referred to the focus of the receiving transducer. However, instead of two-way delay lines, the one-way distance of the synthetic point referred to the transmit transducer focus was used. This was based on the assumption that the super-harmonic signal was generated only near the focus of the transmit transducer, considered the fact that highly focused transducer was used. Thus, in addition to the one-way distance, the focal distance of the transmit transducer was added to the delay. The equation of delay was give below: ∆𝑡 § =𝑧 .) +𝑠𝑔𝑛(𝑧−𝑧 y ) [𝑎𝑏𝑠 𝑧−𝑧 y −𝑟] 𝑐 (6-5) where 𝑧 .) was the transmit focal distance. Results of the VPSA+CFW beamforming for STAR mode super-harmonic imaging at different transmit voltage are shown in Figure 6-11 and Figure 6-12, using phase array and linear array respectively. The enhancement for phase array super- harmonic imaging was not significant. This was because of the small aperture size of the array, which gave a small area of overlapping for synthesis computation. For the linear array imaging, comparing to pre-beamformed images in Figure 5-13, beamformed images were more aggregated. The higher the transmit voltage, better beamforming result was seen. Pseudo-enhancement effect was successfully eliminated 115 in post-beamforming images. Note the linear array experiment was performed in the water instead of tissue background. Figure 6-11: Result of VPSA+CFW method for phase array STAR mode super-harmonic imaging. (A) DAS beamforming, reproduced from Figure 5-11(C). (B) DAS+VPSA beamforming. Table 6-3 summarizes the resolution change from the VPSA beamforming. For the negative change, the resolution was improved. The enhancement on the lateral resolution effect was obvious, where there was a 10% - 50% improvement at different transmit voltages. Due to the blinking nature of MCAs response, the improvement was not uniform. The axial resolution was reduced at some transmit voltages. 116 Figure 6-12: Result of VPSA+CFW method for linear array STAR mode super-harmonic imaging shown in Figure 5-13. (A)- (E) transmit voltage 22, 38, 54, 70, 85 V pp respectively. 25dB dynamic range. 117 Table 6-3: Resolution Improvement for Linear Array Super-harmonic STAR Mode Imaging VPSA+CFW Transmit Voltage 22V 38V 54V 70V 85V -6dB Axial Resolution (mm) 0.326 0.225 0.179 0.251 0.104 Change 51.6% -3% 35.6% 37.1% -38.1% 6dB Axial Resolution (mm) 0.857 1.068 0.333 0.949 0.728 Change -29.8% -19.6% -50.3% 4.3% -7% 6.4 Summary and Discussion In this chapter, the virtual point synthetic aperture beamforming and coherent factor weighting was reviewed and applied to the super-harmonic imaging. The performance of the beamforming was first demonstrated to single element ultrasound bio-microscopy and IVUS imaging, showing improving lateral resolution and SNR. When using VPSA+CFW for dual-element super-harmonic imaging, better SNR was achieved, while the effect of pseudo-enhancement was suppressed. Finally, VPSA beamforming was applied to STAR super-harmonic, where an enhancement of lateral resolution and suppression of pseudo-enhancement was found. 118 Chapter 7 Conclusion and Future Work 7.1 Conclusion In this work, super-harmonics contrast-enhanced imaging, a new method of angiography was studied. The design and analysis for a configurable super-harmonic platform, which included the transducer, electronic system and beamforming method, was discussed. A dual-channel configurable transmit/receive system for super-harmonic imaging was designed, built and tested experimentally. The system was capable of obtaining both B-mode imaging and super-harmonic imaging. Multiple user-defined working modes were selectable, implemented by the FPGA. Pulse generator, low-noise data acquisition unit and configurable signal processing unit were integrated in the system. The system had 56dB SNR and 42dB variable gain. When working with a 6.5MHz, 35MHz dual- frequency transducer, the system was capable of obtaining super-harmonic contrast images of 30dB dynamic range on a vasa vasorum mimicking phantom. The system showed promise for microvascular assessment in preclinical studies. Array transducers and systems were used for the STAR mode super-harmonic imaging. Improvement on the hardware was done to improve the sampling rate. The system was capable of acquiring super-harmonic data. Compared to single element approach, using the array made real-time super-harmonic imaging possible. By using the single focused transducer for excitation, lower power (0.45MPa) threshold was needed to induce super-harmonic, and better SNR (25dB) was acquired. 119 The VPSA beamforming method together with the CFW was studied and applied to the super-harmonic imaging. The beamforming showed promises on the ultrasound microscopy using highly focused transducer and on IVUS imaging using non-focused transducer. When applied to the dual-element super-harmonic imaging, VPSA method enhanced the image SNR and successfully suppressed the pseudo-enhancement. When applied to the STAR mode imaging, the VPSA method was able to enhance the lateral resolution, while suppressed the pseudo-enhancement as well. 7.2 Future Work 7.2.1 Dual-frequency Array with Dynamic Transmit/Receive The resolution of STAR mode super-harmonic imaging are not as good as that of the dual-element super-harmonic. In dual-element mode, the transmitting and receiving transducer are co-focused. While in STAR mode, dynamic receiving focusing with a single transmit focusing is used. Due to the lack of transmit focusing, the resolution of STAR mode super-harmonic images is poorer. Two options are possible: one is to synchronize the motion of the single element transducer with the line acquisition of array receive, to achieve a mechanical dynamic transmit focusing; the other way is to use a dual-frequency array with dynamic transmit and receive focusing ability. A preliminary version of dual-frequency dual-layer array transducer has been presented in the literature [40], in which the low-frequency transmit elements were connected in parallel thus cannot be used for dynamic transmit beamforming. If the sub-element of the low- 120 frequency array can be connected separately, dynamic beamforming will become possible for array-based super-harmonic, giving 100µm theoretical resolution and high imaging speed. 7.2.2 In-vivo Animal Study All the transducers, systems and beamforming methods being studied in the current work were only tested using ex-vivo or in-vitro phantom, limited by the experiment resources. Due to the fact that micro-vascular is very hard to maintain its structure in ex-vivo condition, in-vivo animal studies are necessary. There are two aspects to be considered for in-vivo super-harmonic imaging: I. The acoustic intensity required to induce the MCA non-linear oscillation is high, around 0.4MPa as found in current study. To apply to the animal, thermal and non-thermal effect and their safety limits need to be careful calibrated. One possible way to improve the non-linear signal strength is to develop MCAs with different shell and core composition. II. It is necessary to have an integrated, compact system in order to perform the super- harmonic imaging experiment in pre-clinical setup. The dual-channel super-harmonic imaging system was developed in the current work. For array-based imaging, the system compatibility needs to be improved. The two-stage beamforming method mentioned in Chapter 6 has been successfully implement on FPGA-core hardware [60], which can be the potential direction for the current application. 121 BIBLIOGRAPHY [1] W. Y. Kim, P. G. Danias, M. Stuber, S. D. Flamm, S. Plein, E. Nagel, et al., "Coronary magnetic resonance angiography for the detection of coronary stenoses," N Engl J Med, vol. 345, pp. 1863-9, Dec 27 2001. [2] P. J. Nederkoorn, Y. van der Graaf, and M. G. Hunink, "Duplex ultrasound and magnetic resonance angiography compared with digital subtraction angiography in carotid artery stenosis: a systematic review," Stroke, vol. 34, pp. 1324-32, May 2003. [3] K. Nieman, F. Cademartiri, P. A. Lemos, R. 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Shung, "Development of a real-time, high-frequency ultrasound digital beamformer for high-frequency linear array transducers," IEEE Trans Ultrason Ferroelectr Freq Control, vol. 53, pp. 317-23, Feb 2006. [43] J. Ma, X. Jiang, K. H. Martin, P. A. Dayton, Y. Li, and Q. Zhou, "Dual frequency transducers for intravascular ultrasound super-harmonic imaging and acoustic angiography," in Ultrasonics Symposium (IUS), 2014 IEEE International, 2014, pp. 675-678. [44] G. L. ten Kate, G. G. Renaud, Z. Akkus, S. C. van den Oord, F. J. ten Cate, V. Shamdasani, et al., "Far-wall pseudoenhancement during contrast-enhanced ultrasound of the carotid arteries: clinical description and in vitro reproduction," Ultrasound Med Biol, vol. 38, pp. 593-600, Apr 2012. 124 [45] J. A. Jensen, S. I. Nikolov, K. L. Gammelmark, and M. H. Pedersen, "Synthetic aperture ultrasound imaging," Ultrasonics, vol. 44 Suppl 1, pp. e5-15, Dec 22 2006. [46] J. T. Ylitalo and H. 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Abstract (if available)
Abstract
Super-harmonic contrast-enhanced ultrasound has been demonstrated to be useful for microvascular imaging. By exciting the microbubble contrast agent near its resonance frequency and receiving echo at the super-harmonic, high-contrast high-resolution image from contrast agent could be obtained and discriminated from tissue. ❧ While dedicated dual-element transducers have been designed, no existing integrated system could fully support the application. To fulfill this need, a configurable dual-channel transmit/receive system for super-harmonic imaging was designed, built and tested experimentally. The system was capable of obtaining both B-mode imaging and super-harmonic imaging. Multiple user-defined working modes were selectable, implemented by a field programmable logic gate. The function of pulse generator, low-noise data acquisition unit and configurable signal processing unit were integrated in the system. The system had 56dB SNR and 42dB variable gain. When working with a 6.5MHz, 35MHz dual-frequency transducer, the system was capable of obtaining super-harmonic contrast images of 30dB dynamic range on a vasa vasorum mimicking intravascular phantom. The system showed promise for microvascular assessment in preclinical studies. ❧ Currently, all super-harmonic imaging studies were based on dual-element transducer. Array transducer based imaging, which provides transmit/receive from multiple channels synchronously, could potentially improve the signal coherence and imaging resolution, as well as enhance the speed of imaging. Based on an array beamforming system, with proper software and hardware upgradation, we successfully demonstrated the capability of single-transmit-array-receive mode (STAR) super-harmonics imaging. The system was capable of working with multiple kinds of transducer for the application. ❧ The beamforming method for super-harmonic imaging was studied. Synthetic aperture beamforming, which was a well-accepted method in traditional ultrasound, was used for super-harmonic imaging for the first time. By utilizing the echo information from the superimposed acoustic field of the neighboring emission, an image with higher lateral resolution, higher SNR could be acquired. We demonstrated the method in B-mode ultrasound microscopy, and showed that the method works for super-harmonic imaging as well, in both dual-element and array approaches.
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Li, Yang
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Configurable imaging platform for super-harmonic contrast-enhanced ultrasound imaging
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Viterbi School of Engineering
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Doctor of Philosophy
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Biomedical Engineering
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10/31/2016
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07/20/2016
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beamforming algorithm,cardiovascular disease,contrast-enhanced ultrasound,electronic system,FPGA,intravascular ultrasound,microbubble contrast agent,OAI-PMH Harvest,super-harmonic imaging,ultrasound imaging,ultrasound transducer
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Shung, K. Kirk (
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Tags
beamforming algorithm
cardiovascular disease
contrast-enhanced ultrasound
electronic system
FPGA
intravascular ultrasound
microbubble contrast agent
super-harmonic imaging
ultrasound imaging
ultrasound transducer