Advamces in intravascular ultrasound (IVUS)-based technology

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Content
ADVANCES IN INTRAVASCULAR ULTRASOUND (IVUS)-BASED
TECHNOLOGY

by

Mingyue Yu

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)

May 2018

Copyright 2018 Mingyue Yu
i
Dedication

To my beloved

Parents
Yongfang Sun & Rude Yu

ii
Acknowledgements
First of all, I would like to express my sincere gratitude to my advisors, Dr. K.
Kirk Shung and Dr. Qifa Zhou for their support and guidance. Their mentoring and
encouragement not only guided me through this four-and-half-year Ph.D. journey but
will also inspire me to explore the medical ultrasound imaging in the future. My special
thanks to Dr. Jesse Yen. His instruction gives me a better understanding of
beamforming theory and incites me to investigate more in this field. I would also like to
thank Dr. Mike Shuo-Wei Chen and Dr. Jean-Michel Maarek for serving on my
dissertation and qualification committee and for their indispensable input during my
preparation of this dissertation.
I am grateful to all the help and support from my colleagues and friends inside
and outside the NIH Resource Center for Medical Ultrasonic Transducer Technology,
especially Dr. Ruimin Chen, Dr. Yuling Chen, Dr. Teng Ma, Dr. Yang Li, Dr. Kwok
Ho Lam, Dr. Chi Tat Chiu, Dr. Ying Li, Dr. Zeyu Chen, Mr. Xuejun Qian, Mr. Yang
Lou, Mr. Nestor Cabrera, Mr. Hayong Jung and Mr. Hae Gyun Lim, as well as many
other UTRC members. I am also heartily thankful to collaborators from UC Irvine,
including my project supervisor Dr. Zhongping Chen, my colleagues Ms. Rachel Qu,
Ms. Yan Li, Mr. Buyun Zhang and Mr. Youmin He.
Last but not the least, this dissertation is dedicated to my parents Rude Yu and
Yongfang Sun, my brother Wanchun Yu, my sister Jing Yu and my dear friends for
their unconditional love and unlimited support. They are always there for me through
the ups and downs. Thank you so much for always being on my side during this
challenging but rewarding journey.
iii
Table of Contents
Dedication ............................................................................................................................ i

Acknowledgements ............................................................................................................. ii

List of Tables ...................................................................................................................... v

List of Figures .................................................................................................................... vi

Abstract ............................................................................................................................... x

Chapter 1 Introduction ........................................................................................................ 1

1.1 Clinical Background: Atherosclerosis .............................................................. 1

1.2 Technological Background: Intravascular Ultrasound (IVUS) ........................ 3

1.2.1 Diagnostic IVUS ................................................................................ 5

1.2.2 Therapeutic IVUS .............................................................................. 9

1.3 Motivations and Objectives ............................................................................ 11

1.4 Outline ............................................................................................................. 12

Chapter 2 Intravascular Ultrasound Imaging With Virtual Source Synthetic Aperture
Focusing and Coherence Factor Weighting ...................................................................... 14

2.1 Introduction ..................................................................................................... 14

2.2 Method ............................................................................................................ 20

2.2.1 Virtual Source Synthetic Aperture and Coherence Factor Weighting
................................................................................................................... 20

2.2.2 Simulations in Field II ..................................................................... 25

2.2.3 In Vitro Phantom and Ex Vivo Artery Imaging ................................ 26

2.3 Results ............................................................................................................. 30

2.3.1 Lateral Scan ..................................................................................... 30

2.3.2 Rotational Scan ................................................................................ 32

2.4 Conclusions and Discussions .......................................................................... 37

iv
Chapter 3 Integrated Multi-frequency Intravascular Ultrasound and Intravascular
Photoacoustic Imaging ...................................................................................................... 46

3.1 Introduction ..................................................................................................... 46

3.2 Method ............................................................................................................ 52

3.2.1 Probe design and fabrication ............................................................ 52

3.2.2 Imaging System Setup ..................................................................... 56

3.3 Experimental Imaging Results ........................................................................ 57

3.3.1 Lipid-Mimicking Phantom Imaging ................................................ 58

3.3.2 Ex vivo Human Artery Imaging ....................................................... 59

3.4 Conclusions and Discussions .......................................................................... 61

Chapter 4 Intravascular Ultrasound-based Single Beam Acoustic Tweezer .................... 64

4.1 Introduction ..................................................................................................... 64

4.2 Method ............................................................................................................ 69

4.3 Experimental Results ...................................................................................... 71

4.3.1 Transducer Characterization ............................................................ 71

4.3.2 Force Analysis during Manipulation ............................................... 73

4.3.3 Trapping and Manipulation Results ................................................. 74

4.4 Conclusions and Discussions .......................................................................... 76

Chapter 5 Conclusions and Future Work .......................................................................... 78

5.1 Conclusions ..................................................................................................... 78

5.2 Future Work .................................................................................................... 79

Bibliography ..................................................................................................................... 81

v
List of Tables
Table 4-1: Acoustic Output Exposure Level of SBAT at focus ....................................... 72

vi
List of Figures
Figure 1-1: Pathological progression of atherosclerosis. (adapted from (Libby, 2002)) .... 3

Figure 1-2: Two different types of IVUS catheter: (a) Single-element based Mechanical-
Scanning Catheter; (b) Array-based Electronic-Scanning Catheter. (adapted from
www.bostonscientific.com) ................................................................................................ 5

Figure 1-3: (a) Coronary angiogram, an X-ray image with radio-opaque contrast agent
injected into coronary artery tree (b) cross-sectional IVUS image (c) longitudinal IVUS
image. (adapted from https://www.healthcare.siemens.com/) ............................................ 6

Figure 1-4: Molecule-targeted microbubble conjugated with plasmid DNA for targeted
gene delivery. (adapted from Linsey, 2012) ..................................................................... 11

Figure 2-1: Simulated two-way beam patterns and -6dB beam width contour (black
dashed line) of IVUS transducers with 0.5mm active aperture size: (a) 40MHz IVUS
transducer with flat aperture, (b) 80MHz IVUS transducer with flat aperture, (c) 40MHz
IVUS transducer with aperture spherically focused at 3mm. (d) Simulated beam pattern
from (a) with expected -6dB beamwidth contour (green dash dot line) after virtual source
synthetic aperture and coherence factor weighting (VSSA-CFW) processing. The x-axis
is lateral direction, and the z-axis is axial direction. ......................................................... 16

Figure 2-2: Concept of conventional mechanical scan B-mode imaging, where the
acoustic field of the transducer is shown in dashed line. The red dot denotes the transmit
focus. The green dots denote the imaging points. ............................................................. 20

Figure 2-3: Illustration of IVUS imaging with transducer element undergoing rotational
pullback motion. ............................................................................................................... 21

Figure 2-4: (a) Concept of virtual source synthetic aperture imaging in linear scan mode.
(b) Concept of virtual source synthetic aperture imaging in rotational scan mode. The
number in the image shows the number of overlapped acoustic beams. The red dots
denote the focal point. The green dots denote the imaging point. .................................... 23

Figure 2-5: Schematic diagram of coherence factor weighting (CFW). ........................... 24

Figure 2-6: Schematic diagram of the IVUS imaging system setup. ................................ 28

Figure 2-7: Ultrasound signal processing block diagram. ................................................ 30

Figure 2-8: Simulation results of wire phantom imaging by linear scanning. Dynamic
range: 50dB. ...................................................................................................................... 31

Figure 2-9: Experimental results of wire phantom imaging by linear scanning. Dynamic
range: 50dB. ...................................................................................................................... 32

vii
Figure 2-10: -6dB lateral resolution results of wire phantom images by linear scanning. 32

Figure 2-11: Simulation results of wire phantom imaging by rotational scanning.
Dynamic range: 50dB. Scale bar: 2mm. ........................................................................... 33

Figure 2-12: Experimental results of wire phantom imaging by rotational scanning.
Dynamic range: 50dB. Scale bar: 2mm. ........................................................................... 33

Figure 2-13: -6dB lateral resolution results of wire phantom images by rotational
scanning. ........................................................................................................................... 34

Figure 2-14: Lateral beam profiles of the simulated data for the two different scanning
mode. (a) Linear Scanning; (b) Rotational scanning. Green solid line: Raw data. Red
dashed line: VSSA-Only data. Blue dot dash line: VSSA-CFW data. ............................. 34

Figure 2-15: Experimental imaging result of the two agar-based anechoic cyst phantom
imaging: top row shows phantom with 1mm cyst at different depth; bottom row shows
phantom with different cysts at 3mm. Dynamic range: 50dB. Scale bar: 2mm. .............. 36

Figure 2-16: Contrast-to-noise ratio (CNR) results of the two anechoic cyst phantom
images. (a) CNR of 1mm cysts at different depth. (b) CNR of different cysts at 3mm. .. 36

Figure 2-17: Experimental result of ex vivo human artery imaging: (a) Raw image, (b)
Beamformed image using VSSA-CFW. Dynamic range: 50dB. Scale bar: 1mm. ........... 37

Figure 2-18: Simulated point target image and corresponding coherence factor (CF)
profile. ............................................................................................................................... 39

Figure 2-19: Transducer response: (a) experimental transducer, (b) simulated transducer
in Field II. .......................................................................................................................... 40

Figure 2-20: Simulation results of wire phantom imaging by linear scanning using the
experimental transducer response. Dynamic range: 50dB. ............................................... 41

Figure 2-21: Experimental imaging results of homogeneous tissue-mimicking phantom.
(a) and (b) show the raw and VSSA-CFW images, respectively. Dynamic range: 50dB.
Scale bar: 2mm. ................................................................................................................ 42

Figure 2-22 Simulated anechoic cyst phantom images with NURD. Top row shows Raw
image and bottom row shows VSSA-CFW image: (a) Uniform rotation with 0.225°

angular step size; (b) Rotation with random error angle within (-1.125° -1.125°); (c)
Rotation with sinusoidal modulated error angle within (-0.0675° -0.0675°); (d) Rotation
with kinking within 9°

range. Arrowheads indicate the artifact from NURD. The number
refers to the CNR value for the 2mm cyst. Dynamic range: 70dB. Scale bar: 1mm. ....... 43

Figure 3-1: (a) Schematic of back-to-back integrated IVUS-OCT imaging catheter. (b)
Cross-sectional IVUS image. (c) Cross-sectional OCT image. (d) Fused IVUS-OCT
viii
image. (e) Corresponding hematoxylin and eosin (H&E) histology picture. Scale bar:
0.5mm. (adapted from J. Li et al., 2015) ........................................................................... 48

Figure 3-2: (a) Co-registered NIRS data with IVUS image data in cross-sectional view. (b)
NIRS chemogram acquired during rotational pullback motion. The color-coded map
indicates the probability of the presence of a lipid core (Yellow indicates a high
probability and red indicates a low probability). (c) Co-registered block chemogram with
IVUS image data in longitudinal view. (adapted from: http://www.infraredx.com/) ....... 50

Figure 3-3: (a) Absorption spectrum of water and lipid in near-infrared wavelength range.
There are two optical windows in around 1210 nm and 1720nm where lipid absorption
peaks couple with local minimum of water absorption. (b) Transmission of light through
water. (adapted from (Jansen, Wu, van der Steen, & van Soest, 2014)) .......................... 51

Figure 3-4: Structure of 1-3 composite material. (adpated from http://www.neptune-
sonar.co.uk/) ...................................................................................................................... 53

Figure 3-5: Layer structure of the acoustic stack for the imaging catheter. ..................... 54

Figure 3-6: Photography and schematics of integrated IVUS-IVPA probe of two different
types of design. ................................................................................................................. 56

Figure 3-7: Schematic diagram of the IVUS-IVPA imaging system. ............................. 57

Figure 3-8: Lipid-mimicking phantom imaging. (a) IVUS, (b) IVPA, and (c) combined
IVUS-IVPA images of the phantom by 45MHz PMN-PT probe of Type I; (d) IVUS, (e)
IVPA, and (f) combined IVUS-IVPA images of the phantom by 1-3 composite probe of
Type I; (g) IVUS, (h) IVPA, and (i) combined IVUS-IVPA images of the phantom by
45MHz PMN-PT probe of Type II. .................................................................................. 59

Figure 3-9: Ex vivo human cadaver carotid artery imaging. (a) IVPA image, (b) IVUS
image, and (c) combined IVUS-IVPA image of human carotid artery using 45MHz
PMN-PT Type II probe. .................................................................................................... 60

Figure 3-10: Ex vivo human cadaver coronary artery imaging. (a) 45MHz, (b) 100MHz
and (c) combined multi-frequency IVUS image of human coronary artery. .................... 61

Figure 3-11: Schematic diagram of future probe design: (a) Focused IVUS-IVPA probe;
(b) Fully integrated multi-frequency IVUS-IVPA probe. ................................................. 63

Figure 4-1: Schematic diagram of therapeutic IVUS catheter with sirolimus-loaded
microbubbles. The microbubbles are injected through the imaging catheter and then
displaced onto vessel wall by radiation force from IVUS transducer. (adapted from
Kilroy et al., 2015) ............................................................................................................ 65

Figure 4-2: Trapping model of acoustic vortex: the four-element transducer creates
potential-well that provides trapping force to manipulate microbubbles. (adapted from
(Lo et al., 2015)) ............................................................................................................... 66

ix
Figure 4-3: Experimental setup (a) and acoustic trapping node (b) of surface acoustic
wave based 3D manipulation trapping technique. (adapted from Guo et al., 2016) ......... 67

Figure 4-4: Schematic diagram of single beam acoustic trapping. ................................... 68

Figure 4-5: Experimental configuration for demonstrating intravascular trapping
capability of SBAT. .......................................................................................................... 69

Figure 4-6: Time-domain pulse-echo signal (black solid line) and normalized frequency
spectrum (red dashed line) of the 50MHz transducer. ...................................................... 71

Figure 4-7: (a) The two-dimensional beam profile on the XZ plane with (0,0) position
referring to the focal point of the transducer. (b) The contour plots delineate pressure
level of -3dB, -6dB, -9dB and -12 dB. .............................................................................. 72

Figure 4-8: Illustration of forces exerting on a microparticle (red circle) on tube wall. The
black dashed line denotes the ultrasound beam axis. Gold dashed circle in the tube center
indicates the rotation trajectory of transducer aperture. .................................................... 74

Figure 4-9: Single polystyrene microparticle of 10µμm mean diameter was trapped and
manipulated by the SBAT along mylar film tube. The red dashed circle marks the trapped
microparticle. The green dashed ellipse indicates a reference position to show the motion
of the microparticle. The blue arrow denotes the moving direction of the microparticle. 75

Figure 4-10: A group of polystyrene microparticles of 10µμm mean diameter were trapped
and manipulated by the SBAT along the mylar film tube. The red dashed circle marks the
trapped microparticles. The green dashed circle indicates a reference point to show the
motion of the microparticles. The blue arrow denotes the moving direction of the
aggregated microparticles. ................................................................................................ 76

x
Abstract
Cardiovascular disease (CVD), the number one killer in the United States, causes
around 30.8% of all deaths. Most of CVD is related to a progressive process,
atherosclerosis. Atherosclerotic plaque builds up gradually within the artery wall and may
remain asymptomatic until high-risk plaque (vulnerable plaque) ruptures, resulting in
heart attack or even sudden death. Intravascular ultrasound (IVUS) has frequently been
used as a clinical diagnostic tool for intravascular imaging to measure artery lumen
dimension and the total plaque volume. However, current IVUS technique is not
sufficient for accurate and robust diagnosis of vulnerable plaque, especially regarding
image quality and image information. In addition to the diagnostic benefit, the therapeutic
value of IVUS in the combined use of microbubbles has also been intensively
investigated in recent years. However, a technique for controllable trapping and
manipulation of microscopic targets along artery wall is still needed for efficient therapy
delivery.
In this dissertation study, IVUS technique has been advanced in three aspects: 1.
applied virtual source synthetic aperture (VSSA) focusing and coherence factor
weighting (CFW) to improve IVUS image quality; 2. integrated multi-frequency IVUS
and intravascular photoacoustic (IVPA) imaging to provide both tissue structure and
chemical composition information needed for accurate diagnosis; 3. demonstrated the
trapping and manipulation capability of IVUS-based single beam acoustic tweezer
(SBAT) along curved surface that is geometrically similar to vessel structure. The results
of these studies successfully consolidate the improvements in different aspects of IVUS
technology, which suggests a promising future when the advanced IVUS-based
xi
multimodal intravascular probe will be developed for both robust diagnostic imaging and
efficient therapeutic delivery.

1

Chapter 1 Introduction
1.1 Clinical Background: Atherosclerosis
The clinical motivation of this research study is atherosclerosis, the leading cause
of mortality and morbidity worldwide. It is a chronic disease characterized by the build-
up of atherosclerotic plaque through the accumulation of monocytes, cholesterol, lipids,
fibrous constituents, and various inflammatory cells in the inner lining of the arterial wall.
Most of the cardiovascular diseases (CVD) such as coronary artery disease (CAD),
cerebrovascular accident (CVA) and peripheral arterial disease (PAD) involve
atherosclerosis. Our knowledge of atherosclerosis has changed dramatically over the last
few years. The development of atherosclerotic plaque is outlined in Figure 1-1 (Libby,
2002). The normal artery wall has a three-layer structure: the inner layer called intima
contains a single endothelial cell layer; the middle layer called media is composed of
smooth muscle cell and elastic fiber that enables the dilation of artery wall; the outermost
layer is adventitia that contains fibroelastic tissue. Atherosclerosis is initiated by the
inflammatory response in intima layer triggered by the abnormal accumulation of low-
density lipoprotein (LDL) that carries fat molecules. A snowballing aggregation of lipid
core will happen if high-density lipoprotein (HDL) cannot remove fats effectively from
the endothelial layer. The fibrous cap composed of elastin and collagen forms over the
lipid core as an attempt to heal the atherosclerotic lesion. At an early stage, the
atherosclerotic plaque (atheroma) develops without occupying the lumen area owing to
the artery remodeling, during which the external elastin membrane between media and
2
adventitia expands to keep lumen size and maintain blood flow. The plaque begins to
narrow the lumen once the artery wall can no longer expand outwards. Some of the
atherosclerotic plaque is stabilized with a thick fibrous cap, small lipid core and normal
lumen area. However, some of the plaque becomes vulnerable and prone to rupture. The
rupture of vulnerable plaque triggers thrombus or blood clot formation, which may
partially or totally block the blood flow and cause acute events such as myocardial
infarction or stroke (Lusis, 2000). Healing may also take place after rupture, resulting in a
narrower lumen and a more stable plaque with higher fibrotic composition. The
progression of atherosclerosis is chronic and complicated. Lots of efforts are still needed
to fully understand the atherosclerotic process.
Recent clinical research mainly focuses on the crucial role of vulnerable
atherosclerotic plaque, especially thin-cap fibroatheroma (TCFA) that induces most of
the acute coronary syndrome (ACS) (Kolodgie et al., 2001). TCFA typically has a fibrous
cap thinner than 65um with severe macrophage infiltration and a large lipid core taking
up more than 35% of whole plaque volume under the fibrous cap (Virmani, Burke,
Kolodgie, & Farb, 2003). These distinctive features of TCFA have motivated and guided
the advances in the corresponding diagnostic and therapeutic strategies to effectively
detect and treat atherosclerosis at an early stage before it threatens the patient’s life.
3

Figure 1-1: Pathological progression of atherosclerosis. (adapted from (Libby, 2002))
1.2 Technological Background: Intravascular Ultrasound (IVUS)
Intravascular ultrasound (IVUS) has been widely accepted in clinical practice for
more than 20 years. It is frequently used in conjunction with angiography for
intracoronary imaging to visualize the atherosclerotic plaque and guide stent placement,
which cannot be achieved through angiography. Clinical studies have demonstrated that
IVUS can benefit patients by changing percutaneous coronary intervention (PCI) strategy
and reducing heart attack. Grounded in extensive scientific literature and supportive
clinical outcomes, IVUS will remain valuable and irreplaceable. What is more, it is the
fundamental image modality for many advanced intravascular techniques since it
4
provides a complete image of vessel wall anatomy on which other image information can
be superimposed. In recent years, the therapeutic function of IVUS has also been
increasingly investigated for localized and efficient therapy delivery onto targeted lesion
detected through imaging.
There are two different types of IVUS catheters. One type, mechanical scanning
catheter, contains a single-element transducer mounted on the tip (40~45MHz), as shown
in Figure 1-2(a). A 360° cross-sectional image is acquired by mechanically rotating
imaging catheter through a motor and long torque coil. Another type, electronic scanning
catheter, has a ring-shape phase array transducer (10~20MHz) surrounding the
circumference of the catheter tip, as shown in Figure 1-2(b). Electrical scanning based on
sophisticated transmission/receiving electronics enables 360° imaging of vessel wall
without mechanical rotation. High frame rate (100fps) can be achieved owing to the
electrical scanning, which can avoid the non-uniform rotational distortion (NURD)
inherent to mechanical scan method. However, the resolution and sensitivity of array-
based IVUS are still subpar because of the lower frequency and smaller amount of
elements limited by current fabrication technology. Therefore, mechanical scanning
catheter is always a favorable choice for diagnostic imaging of atherosclerosis. Moreover,
single-element based IVUS holds the promising potential to be advanced to a more
powerful intravascular theranostic tool. In this study, the single-element based IVUS
technology is improved in three different aspects aimed at better diagnosis and treatment
of atherosclerosis.
5

Figure 1-2: Two different types of IVUS catheter: (a) Single-element based Mechanical-
Scanning Catheter; (b) Array-based Electronic-Scanning Catheter. (adapted from
www.bostonscientific.com)
1.2.1 Diagnostic IVUS
Before conducting IVUS imaging, the long (~135cm) imaging catheter is first
advanced into the coronary artery through guidewire (~0.36mm) under the guidance of
the 2D silhouette of the coronary tree provided by angiography as shown in Figure 1-3(a).
During IVUS imaging, the transducer sends ultrasonic pulse towards the artery wall and
then receives echo signal resulting from acoustic impedance mismatch at each radial
position during the rotational pullback motion. The voltage trace (A-line) converted from
echo signal carries structural information of artery wall. The geometrical distance can be
reconstructed using equation 1-1, where d refers to the distance between transducer
aperture and the reflection position, 𝑓
!
is the sampling frequency during data acquisition
(𝑓
!
at least 4 times of the transducer center frequency was chosen in this study), n is
number of sample points, c is sound speed in the propagation medium.

𝑑=
𝑛𝑐
2𝑓
!

1-1
6
The digitized raw A-line data is first passed through a digital band-pass filter to
filter out noise signals outside the transducer bandwidth. Hilbert transform is then
conducted on the filtered A-line data for envelope detection followed by logarithmic
compression to convert voltage magnitude to brightness in dB scale at each pixel.
Multiple A-lines received during rotational pullback motion of catheter can be mapped
into 2D gray-scale cross-sectional image (Figure 1-3(b)) and longitudinal image (Figure
1-3(c)) of a vessel wall. In addition to the essential processing pipeline described above,
smoothing methods such as A-line averaging or median filtering can also be applied to
suppress the random background noise.

Figure 1-3: (a) Coronary angiogram, an X-ray image with radio-opaque contrast agent
injected into coronary artery tree (b) cross-sectional IVUS image (c) longitudinal IVUS
image. (adapted from https://www.healthcare.siemens.com/)

7
Image resolution is the key parameter to evaluate the image quality of different
imaging modalities. The axial resolution and lateral resolution of ultrasound image are
determined by the pulse bandwidth and beamwidth, respectively, as described in equation
1-2 and 1-3 (Zhou, Lau, Wu, & Shung, 2011) , where c is sound speed, 𝑓
!
is center
frequency of the transducer, 𝐵𝑊 is -6 dB fractional frequency bandwidth of the
transducer, 𝐹
#
represents the ratio of focal distance to the aperture dimension of the
transducer, 𝜆 is the wavelength.

𝑅
!"#!$
=
𝑐
2𝑓
!
𝐵𝑊
=
𝜆
2𝐵𝑊

1-2

𝑅
!"#$%"!
=
𝑐𝐹
#
𝑓
!
= 𝜆𝐹
#

1-3
For the single-element IVUS imaging catheter, the natural focal distance 𝑓
!
calculated
using equation 1-4 can be used to evaluate lateral resolution, where D refers to the
diameter of active transducer aperture.

𝑓
!
=
𝐷
!
4𝜆

1-4
According to equations 1-2 and 1-3, a higher center frequency can provide better axial
and lateral resolution. It should be noted that equation 1-3 describes the lateral resolution
at the focal point and the lateral resolution downgrades in the off-focus region. Current
single-element based IVUS catheters with center frequency ranging from 40~45 MHz
can provide 60~100𝑢𝑚 axial resolution and 200~400𝑢𝑚 lateral resolution (Elliott &
Thrush, 1996) (Brezinski et al., 1997). A 60MHz IVUS catheter (Kodama HD IVUS,
ACIST Medical Systems, Inc. Eden Prairie, MN ，United States) is developed to provide
8
a finer axial resolution (<40𝜇𝑚). However, the resolution of a 60MHz transducer is still
not good enough to detect fibrous cap thickness of the vulnerable plaque TCFA.
Imaging penetration depth is another critical image quality index that indicates
how deep the transducer can ‘see’ through the tissue structure, in this case the artery wall,
for the IVUS application. It quantitatively represents the sensitivity of the imaging
system and can be defined as the depth at which signal-to-noise ratio (SNR) falls below
6dB. The ultrasound wave will be attenuated as it travels in biological tissue. The
attenuation as a function of traveling distance z can be quantitatively described by (Shung,
2005):
𝐼(𝑧)= 𝐼
!
𝑒
!!!"
1-5
where 𝐼
!
is the signal intensity at the transducer aperture, 𝛼 is the attenuation coefficient
of the tissue of interest. The attenuation coefficient is frequency dependent and varies a
lot among different tissue types. The penetration depth of a 40~45MHz IVUS catheter is
around 5~8mm. The deep penetration depth enables IVUS imaging to visualize three
layers of a vessel wall, to evaluate plaque burden and to monitor artery remodeling
(Elliott & Thrush, 1996) (Brezinski et al., 1997) (Maresca, Adams, Maresca, & van der
Steen, 2014). Attenuation coefficient increases with higher frequency, which makes the
penetration depth decrease with increasing frequency. Therefore, a trade-off has to be
made between the image resolution and penetration depth.
The contrast of ultrasound image originates from the acoustic impedance
differences within the region of interest. Low soft tissue contrast based on echogenicity
limits the capability of conventional gray-scale IVUS imaging for accurate tissue
characterization of artery wall except for calcified or dense fibrous plaque (Bom, Carlier,
9
van der Steen, & Lancee, 2000). To overcome this limitation, backscattered
radiofrequency (RF) spectrum analysis algorithms for plaque compositional
characterization have been developed and implemented in commercial IVUS imaging
systems used in clinical settings, including: Virtual Histology (VH-IVUS, Volcano
Therapeutics, Rancho Cordova, CA, United States), iMAP (Boston Scientific, Santa
Clara, CA, United States), and integrated backscatter IVUS (IB-IVUS, YD Co., Nara,
Japan) (Nair et al., 2002) (Sathyanarayana, Carlier, Li, & Thomas, 2009) (Kawasaki et al.,
2001). Spectral parameters such as slope, intercept, spectral similarity and integrated
backscatter (IB) are extracted from the normalized spectrum of the RF data within the
region of interest (ROI). Then, the tissue type of each center image pixel within that ROI
is obtained through a classifier constructed from a histology-derived database. A color-
coded map is reconstructed with different plaque constituents labeled with specific colors.
However, the resolution (~100𝑢𝑚 ) is still too low to detect the thin fibrous cap with
thickness smaller than 65𝑢𝑚. Moreover, the reliability of using RF spectrum analysis to
identify necrotic core is also challenged because of the inconsistent and competing
definitions of the necrotic core in the algorithm (Thim et al., 2010). Therefore, a more
robust method for plaque composition characterization is needed to reliably identify
TCFA.
1.2.2 Therapeutic IVUS
Paired with functionalized microbubble contrast agent as shown in Figure 1-4,
IVUS technology has been extended for enhancing localized therapy delivery
10
immediately following the diagnosis during the interventional procedure (Dixon et al.,
2015). Two important mechanisms resulting from the interaction between microbubble
and ultrasonic wave contribute to the improvement in therapy uptake. One is
sonoporation, a mechanism that refers to the transient and reversible pore formation in
cell membrane caused by microbubble vibration under ultrasound stimulation. This
process enhances cell membrane permeability and creates a pass way for small drug
molecules or gene vectors (van Wamel et al., 2006) (Phillips, Klibanov, Wamhoff, &
Hossack, 2012a). Another important mechanism is the acoustic radiation force (ARF)
that arises from the momentum transfer between acoustic wave and microbubble (Dayton,
Allen, & Ferrara, 2002). The ARF exerted on the microbubbles enables localized therapy
delivery by transporting the drug-loaded microbubbles from the blood stream onto the
vessel wall to increase the binding efficiency (J. P. Kilroy, A. L. Klibanov, B. R.
Wamhoff, & J. Hossack, 2012). After identifying the target lesion during the
percutaneous coronary intervention procedure, microbubbles are infused through the
catheter while the IVUS transducer rotates. The displacement ultrasound pulse first
displaces microbubbles from blood flow to the vessel wall. Then the disruption pulse is
sent to disrupt microbubbles for therapeutic agent release and to enhance cell membrane
permeability for therapy uptake. However, merely moving the microbubbles away from
IVUS catheter towards vessel wall cannot ensure manipulation along the vessel wall to
precisely target the diseased location. Manipulating the microscopic objects in a
controllable manner along the vessel wall is still needed to further improve lesion
specificity and adhesion, which is essential for efficient therapy delivery.

11

Figure 1-4: Molecule-targeted microbubble conjugated with plasmid DNA for targeted
gene delivery. (adapted from Linsey, 2012)
1.3 Motivations and Objectives
Based on the clinical and technical background of IVUS technology, IVUS
remains indispensable for intravascular diagnostic imaging and becomes promising for
enhancing intravascular therapy delivery. However several facets still need to be
improved to make it a more powerful tool, which motivates me to advance IVUS
technology in the following three aspects:
(1): As the foundation of intravascular imaging technology both in clinic and
academic research, conventional IVUS at 40~45MHz provides a complete picture of the
artery wall. However, the beam divergence of single element transducer downgrades the
image quality and hinders the further development of this technology. Driven by this, I
present the first studies to apply beamforming method that combines virtual source
synthetic aperture (VSSA) focusing and coherence factor weighting (CFW) to enhance
the IVUS image quality over the whole field of view.
12
(2): With limited resolution and low soft-tissue contrast, IVUS itself cannot
provide accurate tissue characterization information to identify the vulnerable plaque,
which inspires me to integrate complementary imaging techniques that can provide both
better resolution and more reliable tissue composition characterization to overcome this
limitation. I target an integrated imaging catheter with low-frequency IVUS (40~45MHz),
high-frequency IVUS (80~100MHz) and lipid sensitive intravascular photoacoustic
imaging (IVPA) to evaluate plaque vulnerability comprehensively.
(3): Current therapeutic IVUS technology enhances localized therapy only by
transporting the microscopic targets from blood flow onto the surface of the vessel wall.
The controllable manipulation of microparticles along the vessel wall is in great need to
further improve therapy delivery. Therefore ，I developed and tested IVUS-based single
beam acoustic tweezer (SBAT) to solve this issue.
The ultimate goal of this dissertation study is to develop more advanced IVUS
based intravascular technology to accurately detect the vulnerable atherosclerotic lesions
and efficiently deliver therapies to the identified lesion.
1.4 Outline
The dissertation is outlined as follows:
Chapter 1 introduces the clinical and technical background that motivates this
study.
Chapter 2 proposes the application of beamforming method VSSA-CFW to
improve IVUS image quality. This chapter begins with the importance and advantage of
13
applying beamforming method to improve IVUS image quality through the entire field of
view. Then the VSSA-CFW and study method are explained. Results from simulation
and experimental studies are presented to evaluate the performance of this beamforming
method for both lateral and rotational scanning. Image quality in terms of lateral
resolution, contrast-to-noise ratio, and penetration depth is quantitatively assessed and
compared. The results are summarized, and future work is discussed to advance the
proposed method towards clinical application.
Chapter 3 presents the development of integrated multi-frequency IVUS-IVPA
imaging catheters and system. The necessity of multi-modal imaging catheter and the
advantage of the combination of multi-frequency IVUS and IVPA are first introduced.
The multi-modal imaging system is explained. Then two different types of probe design
are presented, and their imaging result of the tissue-mimicking phantom and human
artery are presented. The results are discussed, and future work is outlined.
Chapter 4 investigates the trapping and manipulation capability of IVUS-based
single beam acoustic tweezer along vessel wall. This chapter starts with an introduction
to the significance of this study. The transducer is characterized to evaluate its beam
profile and biosafety parameters. Then the force analysis and experimental results are
presented, which demonstrate the promising capability of the IVUS-based SBAT for
enhancing intravascular therapy delivery. The future study plan is outlined to further
validate the benefit of this technique.
Chapter 5 summarizes the work accomplished in this dissertation and discusses
future endeavors.

14
Chapter 2 Intravascular Ultrasound Imaging With Virtual
Source Synthetic Aperture Focusing and Coherence Factor
Weighting
2.1 Introduction
Intravascular ultrasound (IVUS) has been the most frequently used coronary
artery imaging tool in clinical for measuring lumen dimension and atherosclerotic plaque
volume during percutaneous coronary intervention (PCI) procedure. In addition, IVUS
also plays an important role in the treatment procedure by facilitating stent (type, size)
selection, and evaluating stent deployment to foresee potential later events. In order to
capture more defining characteristics of the atherosclerotic plaque, multimodal
intravascular imaging techniques have been developed. Multi-frequency IVUS imaging
(T. Ma et al., 2015) (J. Ma et al., 2015) and integrated IVUS and optical coherence
tomography (OCT) imaging (J. Li et al., 2014) (J. Li et al., 2015) (Rieber et al., 2006)
were developed to detect thin fibrous cap of thin-cap fibroatheroma (TCFA) utilizing
high resolution capability of high frequency IVUS or OCT. Integrated IVUS and
intravascular photoacoustic (IVPA) imaging (Piao et al., 2015) (P. Wang et al., 2014),
integrated IVUS and near-infrared spectroscopy (NIRS) imaging (Brugaletta & Sabaté,
2014) (Schultz et al., 2010), and integrated IVUS, OCT and Fluorescence imaging (Liang
et al., 2014) (T. Ma, Zhou, Hsiai, & Shung, 2016) were proposed to identify the chemical
composition of the plaque. Integrated IVUS and contrast-enhanced imaging (Y. Li et al.,
2015) (J. Ma, Martin, Dayton, & Jiang, 2014) has also been reported to monitor the
15
proliferation of vasa vasorum. It can be observed that most innovative multimodal
intravascular imaging techniques are based on IVUS for a whole picture of the vessel
wall anatomy. IVUS remains essential and irreplaceable owing to its capability to
evaluate plaque burden and monitor artery remodeling (T. Ma et al., 2016) (Maresca et al.,
2014) (Syed & Hodgson, 2016). As the widely recognized clinical tool for intravascular
imaging and the irreplaceable technical basis of most multimodal intravascular imaging
techniques, however, IVUS image quality beyond natural focus region still requires
further improvement. In current IVUS image, the deeper region in both cross-sectional
and longitudinal view cannot be clearly represented because of the downgraded spatial
resolution, contrast-to-noise ratio (CNR) and signal-to-noise ratio (SNR) resulted from
beam divergence of the single element transducer.
Different approaches have been adopted in order to improve IVUS image quality.
These methods are compared through the simulated pulse-echo beam patterns, illustrated
in Figure 2-1, of IVUS transducer with a 0.5mm active aperture size in PZFlex
(Weidlinger Associates, Cupertino, CA). The two-way beam pattern of a conventional
flat 40MHz IVUS transducer with a natural focus at 1.7 mm is shown in Figure 2-1(a) for
reference. Higher spatial resolution can be achieved by increasing the operational
frequency of IVUS transducer, but with compromised CNR, SNR and penetration depth
of IVUS images (T. Ma et al., 2015). Figure 2-1(b) shows the acoustic radiation pattern
of a flat 80MHz transducer. It provides better resolution at 3-4mm, but the intensity gets
attenuated notably beyond the natural focus, inducing degradation in CNR and SNR.
Mechanical focusing of the IVUS transducer is another method to obtain higher
resolution and SNR within the focal zone (Lee, Jang, & Chang, 2016). The beam pattern
16
of a 40MHz IVUS transducer with mechanical focus at 3mm is illustrated Figure 2-1(c).
Since the mechanical focal point of 3mm is beyond the natural focal point of 1.7mm,
further mechanical focusing will only pull the focal zone towards the transducer surface,
which will in turn decrease the sensitivity and lateral resolution in the rest of field of view
(FOV) (Lee et al., 2016) (Yoon et al., 2015). As presented in Figure 2-1(c), narrow beam
width can be achieved at 1-2mm. Nevertheless, the beam diverges beyond that region,
where more artery structure information resides. What is more, it is practically very
challenging to geometrically focus the transducer with such small aperture size (e.g.
0.5mm×0.5mm).

Figure 2-1: Simulated two-way beam patterns and -6dB beam width contour (black
dashed line) of IVUS transducers with 0.5mm active aperture size: (a) 40MHz IVUS
transducer with flat aperture, (b) 80MHz IVUS transducer with flat aperture, (c) 40MHz
IVUS transducer with aperture spherically focused at 3mm. (d) Simulated beam pattern
17
from (a) with expected -6dB beamwidth contour (green dash dot line) after virtual source
synthetic aperture and coherence factor weighting (VSSA-CFW) processing. The x-axis
is lateral direction, and the z-axis is axial direction.

It can be seen that both increasing the center frequency and mechanical focusing
method could only improve the lateral resolution, CNR, and SNR within a limited region
of FOV with downgraded image quality outside of the focal region. Therefore, these
methods require an accurate alignment between the IVUS catheter and the region of
interest to effectively enhance image quality. However, it is impracticable to accurately
control the exact position of the imaging catheter within arteries of different lumen sizes
in an actual intravascular imaging situation. Accordingly, a feasible approach that can
improve the image quality of conventional IVUS transducer over the entire field of view
would be of important significance in this scenario.
Virtual source concept based synthetic aperture focusing technique was initially
proposed by Passman and Ermert to improve lateral resolution and increase imaging
depth of strongly focused single-element transducer (Passmann & Ermert, 1996). The
virtual source synthetic aperture (VSSA) focusing was then further investigated and
extended to the applications with array transducers (Bae & Jeong, 2000) (Frazier &
O'Brien, 1998) (Nikolov & Jensen, 2002). Synthetic aperture sequential beamforming
(SASB) is one of the well-known beamforming methods that were developed on the basis
of the virtual source concept for array transducers. SASB is composed of two separate
beamformers. In the first beamforming stage, the array transducer transmits and receives
beams at certain focal points. In the second beamforming stage, the transmission-
reception focal points act as virtual sources and make up a virtual array. Based on the
virtual array assumption, delay-and-sum (DAS) algorithm is applied on the focused
18
image lines from the first stage to obtain depth independent lateral resolution. SASB
provides high resolution ultrasound image at a high speed and at the same time massively
reduces data load for different imaging applications (Kortbek, Jensen, & Gammelmark,
2013) (Hansen et al., 2014) (Hemmsen, Rasmussen, & Jensen, 2014) (Di Ianni,
Hemmsen, & Jensen, 2016). Based on the assumption that the transducer focal point can
be viewed as the virtual source of a spherical wave within a certain angular range, VSSA
focusing method synthesizes the image line by combining adjacent beams that
acoustically cover the synthesized line. Owing to the enlarged effective aperture, the
synthesis procedure increases lateral resolution and extend imaging depth (Passmann &
Ermert, 1996) (Liao, Li, & Li, 2004) (Kortbek, Jensen, & Gammelmark, 2007). This
inspires me to apply VSSA method to IVUS imaging with a single-element transducer for
improving lateral resolution and SNR.
Even though the synthesized large aperture from VSSA focusing method helps to
improve resolution and SNR, it at the same time elevates the sidelobe level, which
decreases CNR and results in unwanted artifacts (Frazier & O'Brien, 1998) (Karaman, Li,
& Odonnell, 1995). Amplitude apodization applied on the receive aperture can suppress
the sidelobe level and increase contrast resolution. However, the mainlobe of the beam
pattern gets broadened, and lateral resolution decreased (Frazier & O'Brien, 1998).
Coherence factor calculated from the coherence of delayed RF signals has been used as a
focusing criterion to quantitatively evaluate the focusing quality (Mallart & Fink, 1994)
(Hollman, Rigby, & O'Donnell, 1999). Higher coherence factor value implies that
delayed RF signals are more coherent among each other, or better aligned after delaying.
It has been demonstrated that weighting the delayed beams adaptively using the
19
coherence factor can both reduce the sidelobe level of the VSSA focused radiation
pattern (M. L. Li, Guan, & Li, 2004) and improve the lateral resolution (Liao et al., 2004).
This method is called coherence factor weighting (CFW). The expected -6dB beamwidth
contour after VSSA-CFW processing is illustrated by the green dash dot line in Figure 2-
1(d). The application of this combined method VSSA-CFW to the conventional IVUS
transducer has not yet been investigated.
This motivated me to extend the VSSA-CFW beamforming method to IVUS
application for improving lateral resolution and CNR through the entire FOV. In this
chapter, the applicability of this VSSA-CFW method was first validated through
simulations in Field II (Jensen & Svendsen, 1992). Then the efficacy was explored
experimentally through tungsten wire phantom imaging, homogeneous agar-based
phantom imaging, anechoic cyst phantom imaging and ex vivo human artery imaging.
Both simulated and experimental results of wire target imaging demonstrate
improvements in lateral resolution. Also, CNR improvement is explicitly shown in the
anechoic cyst phantom images. Image depth extension was quantitatively measured
through in vitro imaging of the homogeneous phantom. In the human artery image,
extended image depth, and better layer definition are achieved after image processing. To
the best of my knowledge, this is the first time VSSA-CFW being used for IVUS imaging.
This image processing method can be readily integrated into the imaging system to
improve the image quality of current IVUS imaging catheters. Furthermore, it can refine
the multimodal intravascular imaging techniques to provide more accurate and
comprehensive information for vulnerable plaque identification.
20
2.2 Method
2.2.1 Virtual Source Synthetic Aperture and Coherence Factor Weighting
In conventional ultrasonic brightness mode (B-mode) imaging, a single
transmission is fired to collect the information along the axial direction (or depth
direction) with the underlying assumption that the acoustic beam of the transducer has a
narrow and uniform rectangular shape. As a matter of fact, the narrowest beamwidth can
only be achieved at the focal point; in the out-of-focus region, the resolution is inversely
proportional to the distance from the focal point because of the beam divergence. This
concept is illustrated in Figure 2-2, where the shape of the transducer and beam profile
are delineated using the dashed line.

Figure 2-2: Concept of conventional mechanical scan B-mode imaging, where the
acoustic field of the transducer is shown in dashed line. The red dot denotes the transmit
focus. The green dots denote the imaging points.

In the virtual source synthetic aperture (VSSA) focusing method, the focal point
of the transducer is treated as a virtual source, as shown by the red dot in Figure 2-2. This
virtual source produces a spherical wave within a limited angular range (Frazier &
21
O'Brien, 1998) (Bae & Jeong, 2000). The beam convergence and divergence before and
after the focal point are simplified into two cone shapes, and the spreading angle 𝛼 of
these two cones can be approximately determined by the 𝐹
#
of the transducer using 2-1
and 2-2:

𝛼= 2arctan
1
2𝐹
#
, 2-1
𝐹
#
=
𝑧
!
𝐷
,
2-2

where 𝑧
!
is the focal depth of the transducer and 𝐷 is size of the active aperture (Kortbek
et al., 2007).

Figure 2-3: Illustration of IVUS imaging with transducer element undergoing rotational
pullback motion.

In IVUS imaging with rotating single-element transducer, the transducer element
undergoes rotational pullback motion, as illustrated in Figure 2-3, which combines the
rotational scanning along the azimuthal direction in the cross-sectional view and linear
scanning along the elevational direction in the longitudinal view. The cardiologists count
on the information from both cross-sectional view and longitudinal view to make an
accurate diagnosis. Therefore, in this study, I applied the proposed method in both linear
and rotational scan mode with a single-element IVUS transducer. Two modifications
22
were added to perform VSSA processing for an IVUS transducer. First, the natural focus
was treated as the virtual source. The natural focal depth 𝑓
!
is calculated using 2-3.

𝑓
!
=
𝐷
!
4𝜆
,
2-3
Second, instead of assuming an infinitely small focal point, the -3dB beamwidth region
near the natural focus was considered during VSSA focusing. In the region surrounding
the natural focal point, if the adjacent image point at the same depth is covered within the
-3dB beamwidth region, the corresponding RF signal will be used for aperture synthesis
using 2-4 and 2-5.
For linear scan imaging, mono-static VSSA utilizes the superimposed acoustic
field from scan lines adjacent to each other, and compounds echoes from different
emissions to rebuild the high-resolution image. For an imaging point target in the
overlapping acoustic field, denoted as the green point in Figure 2-4(a), VSSA focusing is
done by delaying and summing the echo signals from adjacent emissions, whose beam
field covers the imaging point, as shown in 2-4. The time delay is determined by the two-
way propagation distance from the virtual source to the imaging point using 2-5. (M. L.
Li et al., 2004).
𝑆
!""#
(t)= 𝑅𝐹 𝑖,𝑡−∆𝑡
!
,
!!!
!!!

2-4

∆𝑡
!
= 2 𝑠𝑔𝑛 𝑧−𝑓
!
𝑎𝑏𝑠 𝑧−𝑓
!
− 𝑟
𝑐
,
2-5

where z is the depth of the imaging point, 𝑓
!
is natural focal depth, r is the distance from
the focal point to the imaging point. N is the number of scan lines used for aperture
synthesis, labeled by the number in Figure 2-4. N is determined by the spreading angle 𝛼,
23
scanning step size, and imaging depth. It should be noted that the delay is negative if the
imaging point is before the focal point (closer to the transducer surface). The assumption
of a virtual source/detector is based on the fact that the acoustic beams from the
transducer surface add coherently at the focus. Therefore the time delay is calculated
relative to the virtual source, instead of the transducer surface (Frazier & O'Brien, 1998)
(M. L. Li et al., 2004). Theoretical deduction of the method is outlined by Passmann and
Ermert in their studies (Passmann & Ermert, 1996).

Figure 2-4: (a) Concept of virtual source synthetic aperture imaging in linear scan mode.
(b) Concept of virtual source synthetic aperture imaging in rotational scan mode. The
number in the image shows the number of overlapped acoustic beams. The red dots
denote the focal point. The green dots denote the imaging point.
24

For rotational scan imaging, the beamforming theory of VSSA is the same as
linear scan, but the beam profile and aperture synthesis have to be considered in the polar
coordinates instead of the Cartesian coordinates. As shown in Figure 2-4(b), during the
rotational scanning, the position of the transducer surface changes along a circular path.
The radius of the circular path is determined by the size of the imaging probe. The
angular step size during the rotation, beam spreading angle, and imaging depth, together
determine the number of scan lines that can be used for beam synthesis, as represented by
N in 2-4. Different from the linear scan situation, the beam overlap decreases with
increasing depth during the rotational scan. Accordingly, N is not balanced before and
after the focal point because of the beam separation during rotation. The received RF data
was first passed through a 4
th
order Butterworth filter with a passband of 20-60MHz.
Then VSSA focusing was done on the filtered data using 2-4 and 2-5. The VSSA-Only
image was formed using 𝑆
!""#
in 2-4 through envelope detection and logarithmic
compression.

Figure 2-5: Schematic diagram of coherence factor weighting (CFW).

25
In addition to the VSSA, coherence factor weighting (CFW) was applied to the
synthesized beam 𝑆
!""#
in 2-4 to further improve focusing quality and contrast resolution
(M. L. Li et al., 2004) (P. C. Li & Li, 2003) (Liao et al., 2004). The coherence factor was
calculated based on the signal coherence of delayed RF signals involved in the
synthesized beam as indicated in 2-6. In the equation 2-6, the denominator represents the
total energy (coherent and incoherent) of the delayed RF signals included for aperture
synthesis. The numerator represents the coherent energy portion. The coherence factor
equals 1 when the delayed RF signals are perfectly coherent. It decreases if focusing error
occurs and falls to 0 in the case of zero-mean random noise, or when the delayed signals
are completely misaligned with each other (Figure 2-5). Therefore, the CFW method is
able to suppress the sidelobes arising from VSSA focusing while further narrowing the
mainlobe. After the VSSA-CFW processing, the VSSA-CFW image was formed based
on 𝑆
!""#!!"#
in 2-7 after envelope detection and logarithmic compression.
Wt
!"
(t)=
RF(i,t−∆t
!
)
!!!
!!!
!
N |RF(i,t−∆t
!
)|
! !!!
!!!

2-6
𝑆
!""#!!"#
𝑡 = 𝑆
!""#
Wt
!"
(t) 2-7
2.2.2 Simulations in Field II
In order to assess the efficacy of the proposed method for IVUS imaging,
simulations were conducted using Field II, an acoustic simulation tool based on Matlab
(Jensen & Svendsen, 1992). Both linear and rotational scan imaging were simulated to
comprehensively investigate the performance of the method for IVUS imaging. I first
defined a 40MHz IVUS transducer with a flat 0.5mm×0.5mm aperture to match the
26
transducer that I used in the experiments. Then two types of phantoms were defined: five
point targets with 1mm axial separation and 0.25mm lateral separation were defined for
lateral scan simulation; five point targets with 1mm axial separation and 60° angular
separation counterclockwise were defined for rotational scan simulation. RF signals were
acquired by updating phantom locations according to the linear or rotational scanning
trajectories. The step size is 3.6 𝜇𝑚 and 0.36° for the linear scan and rotational scan,
respectively. VSSA focusing and CFW method mentioned above were applied to the
simulated RF signals. After processing, VSSA-Only and VSSA-CFW images were
formed by envelope detection and logarithmic compression. Lateral resolution at
different depths of Raw, VSSA-Only, VSSA-CFW images was measured to quantify the
effect of the proposed method on the lateral resolution. Image resolutions were
determined by the -6dB beamwidth of the envelope signals from wire targets at different
depths.
2.2.3 In Vitro Phantom and Ex Vivo Artery Imaging
Imaging data were acquired from phantoms and ex vivo human artery to validate
the performance of VSSA-CFW. I fabricated two wire phantoms of different distribution
geometries (linearly separated and circularly separated) using five pieces of 10𝜇𝑚
tungsten wire to study the resolution improvement for the linear scan and the rotational
scan. In the linearly separated wire phantom, the five pieces of wire are separated by
0.2mm and 1mm along lateral and axial direction, respectively. In the circularly separated
27
wire phantom, the wires are separated by 60° and 1mm along angular (counterclockwise)
and axial direction, respectively. Wire phantom images were acquired by linearly and
rotationally scanning the IVUS transducer within deionized water. The lateral and axial
resolutions were measured in the same way as that for simulated images. A cylindrical
agar based tissue-mimicking phantom with 3mm diameter lumen in the center was
fabricated to investigate the improvement in imaging depth. During imaging, the central
lumen was filled with deionized water. An image frame of the homogeneous tissue-
mimicking phantom was captured. The noise image frame was also captured by turning
off the pulser/receiver. The acquired phantom image and noise image data first passed
through envelope detection and logarithmic compression. Then, the signal and noise
mean images were obtained through low-pass filtering. Image signal-to-noise ratio (SNR)
was calculated from the difference between signal and noise mean image. Imaging depth
was determined as the depth where SNR of the image falls below 6dB. For ex vivo
imaging, a section of human cadaver artery sample was fixed coaxially using congealed
agar solution within a 30mL beaker. During imaging, the beaker was mounted on a rotary
motor to obtain the cross-sectional image data.
To evaluate the improvements in CNR quantitatively, two cylindrical agar-based
anechoic cyst phantoms with 3mm diameter lumen in the center were fabricated. One
holds three cysts of 1mm in diameter located at 2mm, 3mm and 4mm away from lumen
center. The other phantom has three cysts of 0.5mm, 1mm and 2mm in diameter located
at 2mm away from the center of the lumen (Turnbull, Lum, Kerr, & Foster, 1992).
During imaging, both the central lumen and cyst lumens were filled with deionized water.
The CNR was computed for every cyst within the phantom using a 50dB dynamic range.
28
CNR was calculated using equation 2-8 (C. Kim et al., 2013) (Moghimirad, Hoyos,
Mahloojifar, Asl, & Jensen, 2016) :

𝐶𝑁𝑅=
|𝜇
!
−𝜇
!
|
(𝜎
!
!
+𝜎
!
!
)
,
2-8
where 𝜇
!
and 𝜇
!
is the mean intensity values (in decibels) within the cyst region and
background region, respectively. 𝜎
!
and 𝜎
!
is the standard deviations of cyst and
background regions, respectively. In this situation, the background region is the tissue-
mimicking agar phantom region.

Figure 2-6: Schematic diagram of the IVUS imaging system setup.

A 40MHz side-looking miniature IVUS transducer of 1 mm in outer diameter and
0.5mm×0.5mm aperture size was used to acquire images. The simulated beam profile of
the transducer is shown in Figure 2-1(a), where the narrowest beamwidth is achieved near
29
the natural focus (1.7mm). Beam diverges at the off-focus region. During linear scanning,
the transducer was mounted on a linear stage with the transducer aperture facing towards
the wires within the water bath. 1000 scan lines were acquired with a 3.6𝜇𝑚 step size to
form the wire phantom image. During rotational scanning, the transducer tip remained
stationary within the water-filled lumen of the imaged sample (wire phantom, anechoic
cyst phantom, and human artery) and cross-sectional images were achieved by rotating
the sample using an integrated servo motor (Moog Animatics, Milpitas, CA, USA) (X. Li
et al., 2011a). 1000 scan lines were acquired with a 0.36° angular step size corresponding
to around 𝜆/12 step size at the transducer surface to ensure enough beam overlap
between adjacent scanlines. The imaging system setup is displayed in Figure 2-6. A JSR
DPR500 pulser/receiver was used to drive the transducer, receive, and amplify the RF
signals. The amplified RF signals were digitized by a 12-bit digitizer (Gage Applied
Technologies, Lockport, IL) at a sampling rate of 500MHz. The motor rotation and RF
data acquisition were synchronized by a function generator to ensure acquiring 1000 scan
lines per rotation of the motor. A custom-developed LabView program (National
Instruments, Austin, TX) was used to control the imaging procedure and save the raw RF
data for offline processing and image display. The saved raw RF data were processed
using different pipelines as shown in Figure 2-7 to get raw, VSSA-Only and VSSA-CFW
images.
30

Figure 2-7: Ultrasound signal processing block diagram.
2.3 Results
The in vitro phantom imaging and ex vivo human artery imaging results
demonstrate that the beamforming method VSSA-CFW can improve the image quality of
single element-based IVUS imaging by enhancing the lateral resolution, extending
imaging depth, and increasing CNR.
2.3.1 Lateral Scan
Figure 2-8 and Figure 2-9 show images acquired by a linear scan in the simulation
and the experiment, respectively. In both the simulation results and experimental results,
there is obvious resolution enhancement in the beamformed images across the whole
imaging range, compared to the raw images. VSSA focusing decreases the mainlobe
width and therefore an almost depth-independent lateral resolution is obtained, as
31
indicated in Figure 2-10. But the sidelobe level becomes higher, as shown in the VSSA-
Only images in Figure 2-8 and Figure 2-9. After applying CFW onto the VSSA focused
beam, the sidelobes are effectively suppressed, and lateral resolution gets further
improved. The lateral resolution is enhanced from 165-524µμm to 126-143µμm after
VSSA-CFW processing. The beam profile of wire phantom is plotted in Figure 2-14(a) to
better illustrate changes throughout the image processing procedure. Mainlobe width is
narrowed after VSSA focusing, but accompanied with more obvious sidelobe. After
VSSA-CFW processing, the side lobe is lowered, and main lobe width is further
narrowed. For instance, sidelobe level is reduced by 34dB and resolution increased by
97µμm for the simulated point target at 3.8mm after applying the CFW.

Figure 2-8: Simulation results of wire phantom imaging by linear scanning. Dynamic
range: 50dB.

32

Figure 2-9: Experimental results of wire phantom imaging by linear scanning. Dynamic
range: 50dB.

Figure 2-10: -6dB lateral resolution results of wire phantom images by linear scanning.
2.3.2 Rotational Scan
The simulation results in Figure 2-11 and experiment results in Figure 2-12 of
rotational scanning also show an increase in the lateral resolution over the whole imaged
range. The resolution results are summarized in Figure 2-13. Different from the linear
33
scan mode, the VSSA focusing alone does not improve the resolution of rotational scan
mode since the moderate expansion of the synthesized aperture cannot compensate the
beam divergence. However, after applying CFW, the lateral resolutions increase by up to
42% for both simulation and experimental results. As shown in Figure 2-14(b), the beam
profile of the wire phantom clearly shows the improvement after VSSA-CFW processing.
For instance, 11dB sidelobe suppression and 2.5° resolution increase were obtained for
the simulated point target at 3.5mm after CFW. Similar to the linear scan case, the axial
resolution maintains after applying the VSSA-CFW.

Figure 2-11: Simulation results of wire phantom imaging by rotational scanning.
Dynamic range: 50dB. Scale bar: 2mm.

Figure 2-12: Experimental results of wire phantom imaging by rotational scanning.
Dynamic range: 50dB. Scale bar: 2mm.
34

Figure 2-13: -6dB lateral resolution results of wire phantom images by rotational
scanning.

Figure 2-14: Lateral beam profiles of the simulated data for the two different scanning
mode. (a) Linear Scanning; (b) Rotational scanning. Green solid line: Raw data. Red
dashed line: VSSA-Only data. Blue dot dash line: VSSA-CFW data.
35
From Figure 2-14, it can also be found that the noise level is suppressed while the
resolution is improved. This is because of the noise-suppressing nature of the VSSA-
CFW beamforming method which is capable of removing random non-coherent signals
while enhancing the coherent signals. Owing to the decrease in noise floor, the imaging
depth was improved from 4.7mm to 6.5mm after applying VSSA-CFW through the in
vitro measurement on the agar-based tissue-mimicking phantom.
The experimental result of agar-based anechoic cyst phantom imaging is shown in
Figure 2-15. The calculated CNR values are summarized in Figure 2-16. The CNR
improves after applying VSSA. However, the higher side lobe still creates artifacts within
the cyst region. After applying CFW onto the VSSA focused image, the CNR further
improves, which demonstrates that CFW can effectively suppress the sidelobes and result
in a higher CNR for the cyst of different sizes and at different depths. In the VSSA-CFW
image, superior CNR can be observed through the more distinct boundary between
anechoic area and surrounding agar phantom, as well as at the boundary of the central
lumen.
36

Figure 2-15: Experimental imaging result of the two agar-based anechoic cyst phantom
imaging: top row shows phantom with 1mm cyst at different depth; bottom row shows
phantom with different cysts at 3mm. Dynamic range: 50dB. Scale bar: 2mm.

Figure 2-16: Contrast-to-noise ratio (CNR) results of the two anechoic cyst phantom
images. (a) CNR of 1mm cysts at different depth. (b) CNR of different cysts at 3mm.

37
IVUS images of the human artery are shown in Figure 2-17. The suppressed noise
floor and enhanced resolution help to better delineate media layer structure of the artery
wall, as indicated by the yellow arrowheads.

Figure 2-17: Experimental result of ex vivo human artery imaging: (a) Raw image, (b)
Beamformed image using VSSA-CFW. Dynamic range: 50dB. Scale bar: 1mm.
2.4 Conclusions and Discussions
This study demonstrates that the beamforming method of combining virtual
source synthetic aperture and coherence factor weighting can effectively improve the
overall quality of IVUS images, including lateral resolution, imaging depth, and CNR.
This is for the first time, to the best of my knowledge, that the VSSA-CFW is applied to
single element-based IVUS imaging.
In this application, the natural focal point of the transducer is treated as a virtual
source based on which the synthetic aperture processing is performed. In the linear scan
mode, VSSA focusing improves the lateral resolution by synthesizing a large aperture.
However, this aperture synthesis procedure brings in high sidelobe level. CFW is
38
therefore needed to reduce the sidelobe level since it can modulate the image intensity
according to the coherence of the delayed RF signals. Coherence factor is derived from
the spatial spectrum of the delayed RF signals. It serves as a quantitative focusing
criterion to weight the synthesized beams according to the coherence of the delayed
scanlines. For a point target, the coherence factor is highest along the synthesized beam
axis that passes through the point target, since the scan lines are highly coherent after
focusing delay. However, for the synthesized beam that’s off the point target, the
coherence factor will decrease because of the steering error for focusing delay calculation.
Coherence factor peaks at the center of point target (center of the mainlobe) and
decreases with increasing distance to the center of the point target. Therefore, the image
intensity from directions other than the central synthesized beam will be suppressed. The
simulated image and corresponding coherence factor profile of the point target located at
3.8mm in linear scan mode were plotted in Figure 2-18 to better illustrate the function of
CFW on narrowing main lobe and suppressing side lobe. In the rotational scan mode,
VSSA focusing alone cannot improve the resolution because of the rotation geometry.
The synthesized aperture size increases with the depth under linear scan mode. But under
rotational scan mode, the synthesized aperture size is not large enough to compensate the
broadening of the beam pattern. As imaging depth increases, the synthesized aperture
only experiences moderate expansion (Kortbek et al., 2007) (Andresen, Nikolov,
Pedersen, Buckton, & Jensen, 2010). However, after applying CFW, the lateral resolution
and CNR still get obvious improvement. Therefore, VSSA-CFW can improve lateral
resolution and CNR for both linear and rotational scan at all depths beyond the natural
focal region, where the most meaningful clinical information potentially lies.
39

Figure 2-18: Simulated point target image and corresponding coherence factor (CF)
profile.

The axial resolution for both simulation and experimental images maintains after
processing. Different from the simulated results, a cosmic tail artifact can be observed in
experimental VSSA-Only and VSSA-CFW image in Figure 2-9. This artifact originates
from the transducer ringing tail response as shown in Figure 2-19, which is an inherent
artifact of the transducer. During VSSA focusing, the sum of low-resolution lines
(unfocused A-line in RAW image) composes the high-resolution line (beamformed line
in VSSA-Only and VSSA-CFW image). For the transducer response with a longer
ringing tail as shown in Figure 2-19(a), the amplitude of those low-resolution lines
decreases gradually and smoothly along the axial direction (Kortbek et al., 2013).
Therefore, the resulting high-resolution lines will contain the oscillating signal, which
appears like the cosmic tail. In comparison, a short and clean transducer response as
shown in Figure 2-19(b) was used in the Field II simulation. The amplitude of the high-
resolution line changes more sharply along axial direction since there are less periods in
the low resolution lines. Accordingly, in this simulation case, VSSA-Only and VSSA-
CFW images didn’t suffer from the cosmic tail artifact. To verify the origin of the artifact
40
and demonstrate the effect of the oscillating tail, I simulated the images by replacing the
response in Figure 2-19(b) with the measured transducer response in Figure 2-19(a).
Results of the simulation are presented in Figure 2-20, where the cosmic tail can also be
observed in the VSSA-Only and VSSA-CFW images. The cosmic artifact issue can be
potentially solved by improving the transducer property. The ringing tail in a transducer
response is mainly related to the backing layer of the acoustic stack. By optimizing the
material selection, the thickness of the backing layer and also the bonding between the
backing layer and the piezoelectric layer, the tail signal in the transducer response can be
minimized. The application of VSSA-CFW method can effectively improve the lateral
resolution and contrast-to-noise ratio. But the image quality along axial direction mainly
relies on the transducer property. Therefore, the proposed VSSA-CFW method, together
with the modified transducer, is capable of providing IVUS images with even better
quality.

Figure 2-19: Transducer response: (a) experimental transducer, (b) simulated transducer
in Field II.

41

Figure 2-20: Simulation results of wire phantom imaging by linear scanning using the
experimental transducer response. Dynamic range: 50dB.

The summation and weighting procedure of VSSA-CFW inherently increases the
image signal-to-noise ratio (SNR) and consequently extends the imaging depth. As
indicated by the red circle in Figure 2-21, the imaging depth for the raw image and
VSSA-CFW image is 4.7mm and 6.5mm, respectively. A 1.8mm imaging depth
extension was obtained for the agar-based phantom. Higher image SNR also enables the
adoption of a higher frequency transducer to achieve higher resolution without sacrificing
the penetration depth.
42

Figure 2-21: Experimental imaging results of homogeneous tissue-mimicking phantom.
(a) and (b) show the raw and VSSA-CFW images, respectively. Dynamic range: 50dB.
Scale bar: 2mm.

The VSSA-CFW beamforming method holds a promising potential for clinical
translation. It can be integrated into the current commercial IVUS imaging system readily
to post-process the acquired IVUS imaging data. Several future works need to be done to
translate this technique to the clinic. First of all, non-uniform rotation distortion (NURD)
is one of the main issues in current IVUS image. It is of great importance to evaluate the
effect of NURD on VSSA-CFW method to further validate the efficacy of this method
for the clinical setting. To study the effect of NURD on the performance of VSSA-CFW
method, I run the simulation for anechoic cyst phantom imaging with different kinds of
NURD. The simulated cyst phantom holds three cyst targets within a 90
o
region: 1mm,
2mm and 3mm in diameter located at 3mm, 5mm, and 7mm, respectively. Three different
types of NURD were simulated: type I is caused by random error angle; type II is caused
by sinusoidal modulation of the rotation which creates a periodic error angle; type III is
the extreme example of NURD (‘kinking’), when the catheter gets stuck and then shoot
43
through. Different types of NURD bring in different artifacts in the raw image as
indicated by arrowheads in the top row of Figure 2-22. IVUS images with type III NURD
typically cannot be used for further diagnostic analysis. The image generated by uniform
rotation was also simulated to compare with images with the non-uniform distortion
effect. With uniform rotation, 400 scan lines were created within 90° region with 0.225°
angular step size.

Figure 2-22 Simulated anechoic cyst phantom images with NURD. Top row shows Raw
image and bottom row shows VSSA-CFW image: (a) Uniform rotation with 0.225°

angular step size; (b) Rotation with random error angle within (-1.125° -1.125°); (c)
Rotation with sinusoidal modulated error angle within (-0.0675° -0.0675°); (d) Rotation
with kinking within 9°

range. Arrowheads indicate the artifact from NURD. The number
refers to the CNR value for the 2mm cyst. Dynamic range: 70dB. Scale bar: 1mm.

As shown in column (b) of Figure 2-22, the VSSA-CFW performs well with small
random error angle such as 1.125°. The jittering artifact was indicated by the red
arrowhead in the raw image under 1.125°

error

angle. With small error angle, the
synthetic aperture focusing method still works since the acquired RF signals are still
coherent. The following CFW is capable of compensating the steering error and further
44
improving the image quality. Similar to the random error angle case, VSSA-CFW works
well under small sinusoidal error angle as shown in column (c) of Figure 2-22. Within the
kinking region, as indicated by the blue arrowhead in column (d) of Figure 2-22, the
image failed to represent any real structure information. Outside the kinking region,
VSSA-CFW is still effective to enhance the image quality as shown by the 2mm and
3mm cysts. CNR for the 2mm cyst is calculated and listed in Figure 2-22 to illustrate the
improvement. Contrast-to-noise ratio still gets improved under moderate NURD artifact.
VSSA-CFW can still perform well under moderate random or periodic error, and outside
the kinking region. Under severe NURD, the artifact can be clearly detected and the
images would not be used for further diagnosis.
The performance of VSSA-CFW on in vivo imaging data needs to be investigated
to further validate this technique. Besides the NURD, the motion of blood vessel and
imaging catheter during in vivo imaging can affect the performance of VSSA-CFW. The
synthetic aperture method is subject to motion artifact. Unwanted image artifacts can
arise if the movement of the imaged target is more than λ/4 during the period needed to
receive all the scanlines included in aperture synthesis (O'Donnell & Thomas, 1992). By
increasing the imaging speed (higher pulse repetition frequency and higher scanning
speed) and adopting motion compensation procedures, the motion artifact problem is
supposed to be solved and will not degrade the performance significantly. VSSA-CFW
method also needs to be validated on the 3D imaging data acquired under 3D rotational
pullback motion, during which the acquired RF signal corresponds to both rotational and
linear position shift. Last but not least, statistical studies about the improvement of the
45
method on different types of plaques need to be performed to further validate the efficacy
of the image processing method.
In summary, VSSA-CFW can effectively improve image quality in both cross-
sectional and longitudinal planes beyond the natural focal point, which ensures the full
coverage of the region of interest for vessels of different sizes. What’s more, it can also
be applied to intravascular photoacoustic imaging, for which lower noise level and better
lateral resolution are highly desired (Liao et al., 2004). The improvement in image quality
will also be a benefit for those increasingly miniaturized IVUS transducers, which is
indispensable for advancing current multimodal intravascular imaging techniques. IVUS
images with improved lateral resolution, extended imaging depth, and enhanced CNR in
both cross-sectional and longitudinal view can assist cardiologists to acquire more
accurate plaque volume measurement, layer structure delineation, and artery remodeling
evaluation.

46
Chapter 3 Integrated Multi-frequency Intravascular
Ultrasound and Intravascular Photoacoustic Imaging
3.1 Introduction
According to current histopathological studies, thin-cap fiberoatheroma (TCFA),
the most crucial vulnerable plaque, has been associated with distinctive clinical markers
regarding both morphological features (thickness, size) and tissue compositions (such as
fibrous cap, lipid-rich necrotic core, calcification, macrophages). These findings have
raised an unmet need for a novel intravascular imaging technology with high resolution
and robust composition-specific information superimposed on the complete picture of
entire plaque volume provided by conventional intravascular ultrasound (IVUS) imaging
at 40 to 45MHz frequency range. However, this information cannot be obtained through
only one imaging modality because each intravascular imaging modality has a unique
view of vulnerability information accompanied by inherent limitations not seen in other
modalities. In this context, a synergistic multi-modal intravascular technique comes onto
the center of the stage (T. Ma et al., 2016).
As discussed in Chapter 2, the beamforming method VSSA-CFW can effectively
improve the image quality of conventional IVUS. Yet, the resolution is still insufficient
to detect the thickness of the thin fibrous cap. Integrated intravascular ultrasound and
optical coherence tomography (OCT) imaging (Figure 3-1) has been developed to
combine the deep penetration depth (5~10mm) of IVUS and the high resolution of OCT
(10~30𝜇𝑚) to provide higher sensitivity and specificity for detecting TCFA (J. Li et al.,
47
2014) (Bourantas et al., 2013) (J. Li et al., 2015). As the optical analog of gray-scale
ultrasound imaging, OCT measures the time delay of back-scattered infrared light using
low-coherence interferometry technique to reconstruct the microstructural information of
blood vessel wall. Unfortunately, the fabrication cost of the optic-ultrasound hybrid
imaging system is very high, which challenges the translation procedure and
compromises the long-term clinical utility of this technique (Maresca et al., 2014).
Inspired by integrated IVUS-OCT technology, multi-frequency IVUS imaging with
conventional frequency IVUS transducer and ultrahigh frequency transducer has proven a
relatively simple and cost-effective substitute for IVUS-OCT system to overcome the
resolution limit of conventional IVUS imaging (T. Ma et al., 2015). The low-frequency
transducer (40~45MHz) can visualize the total plaque burden and monitor artery
remodeling owing to its accurate and deep penetrating capability (Puri, Worthley, &
Nicholls, 2011). And the ultrahigh-frequency (80~150MHz) with improved resolution
holds the potential to measure the thickness of thin fibrous cap (X. Li et al., 2011b) (T.
Ma et al., 2015).
48

Figure 3-1: (a) Schematic of back-to-back integrated IVUS-OCT imaging catheter. (b)
Cross-sectional IVUS image. (c) Cross-sectional OCT image. (d) Fused IVUS-OCT
image. (e) Corresponding hematoxylin and eosin (H&E) histology picture. Scale bar:
0.5mm. (adapted from J. Li et al., 2015)

In recent years, multiple intravascular techniques have been investigated to
analyze chemical composition, especially lipid content, of atherosclerotic plaque for
tissue characterization. Most of these techniques make use of the specific spectrum in
near-infrared (NIR) range (400~2400nm) to differentiate tissue types (Moreno et al.,
2002). Intravascular near-infrared spectroscopy is the first intravascular technique to
identify lipid composition within plaque based on the reflectance spectra of the artery
wall (Moreno et al., 2002) (Waxman et al., 2009). The first commercial dual-modality
IVUS-NIRS intravascular catheter and its accompanying intravascular imaging system
(TVC Imaging System) have been developed. (Infraredx, Inc., Burlington,
Massachusetts). Multiple spectral data are collected and chemical measurements are
performed along the artery segment during the rotational pullback motion of IVUS
49
imaging. The probability of the presence of lipid core plaque (LCP) is then calculated
through the validated predictive algorithm and mapped to 2D color-coded chemogram as
shown in Figure 3-2 (a) and Figure 3-2(b). The four-color block chemogram is also
displayed for every 2-mm chemogram segment (Figure 3-2 (c)). In addition, lipid core
burden index (LCBI) is computed as a quantitative index of the presence of LCP in the
interrogated artery segment. As shown in Figure 3-2, the anatomical IVUS image data
and compositional NIRS spectral data are co-registered to provide comprehensive
detection of vulnerable atherosclerotic plaque. NIRS displays lipid distribution along the
lateral or transversal direction. However, NIRS lacks depth-resolved spatial resolution
and therefore cannot delineate the size and location of the lipid-rich necrotic core. Aside
from NIRS, optical intravascular techniques such as Near-infrared fluorescence (NIRF)
imaging (Yoo et al., 2011) ，Raman Spectroscopy (Buschman et al., 2000) (van de Poll,
Romer, Puppels, & van der Laarse, 2002) ，and fluorescence spectroscopic imaging
(Stephens, Park, Sun, Papaioannou, & Marcu, 2009) ，show great potential to provide
chemical-specific image contrast. Nonetheless, they all suffer from the lack of depth-
resolution needed for cross-sectional mapping of the detected lipid deposition.
50

Figure 3-2: (a) Co-registered NIRS data with IVUS image data in cross-sectional view. (b)
NIRS chemogram acquired during rotational pullback motion. The color-coded map
indicates the probability of the presence of a lipid core (Yellow indicates a high
probability and red indicates a low probability). (c) Co-registered block chemogram with
IVUS image data in longitudinal view. (adapted from: http://www.infraredx.com/)

Intravascular photoacoustic (IVPA) imaging is a hybrid imaging modality that
measures optical absorption contrast with ultrasonic resolution. It detects ultrasound
waves generated by the artery wall in response to the pulsed laser light excitation. IVPA
is a logical addition to IVUS since they share the same ultrasonic transducer, a
relationship which facilitates the integration of IVPA and IVUS in terms of both imaging
catheter and imaging system. The vibrational overtone absorption spectrum of C-H
chemical bonds in the 1.2𝜇𝑚 and 1.7𝜇𝑚 spectral band provides a new window (Figure
3-3) where the optical absorption peaks of lipid couple with a local minimum of water
absorption. Accordingly, the lipid-specific photoacoustic signal can be excited from the
atheromatous plaque and a depth-resolved photoacoustic image reconstructed by
selecting the excitation wavelength within 1.2𝜇𝑚 or 1.7𝜇𝑚 spectral band (Jansen, van
51
der Steen, van Beusekom, Oosterhuis, & van Soest, 2011) (H. W. Wang et al., 2011) (P.
Wang, Wang, Sturek, & Cheng, 2012). IVPA imaging at 1.7𝜇𝑚 spectral band is gaining
more attention for its application in lipid plaque detection since the high optical
absorption of the first overtone of C-H bond enhances the amplitude of the photoacoustic
signal. Lipid-laden plaque can be detected even within the blood at 1.7𝜇𝑚 spectral band
(B. Wang et al., 2012) (P. Wang et al., 2012). An IVPA imaging system at 1.7𝜇𝑚 with
one frame per second frame rate has been developed by our group in collaboration with
the optical research group from University of California, Irvine (UCI) to quantitatively
detect and map the lipid core within atherosclerotic plaque (Piao et al., 2015).

Figure 3-3: (a) Absorption spectrum of water and lipid in near-infrared wavelength range.
There are two optical windows in around 1210 nm and 1720nm where lipid absorption
peaks couple with local minimum of water absorption. (b) Transmission of light through
water. (adapted from (Jansen, Wu, van der Steen, & van Soest, 2014))

Based on the progress in the multi-frequency IVUS and lipid-sensitive IVPA, the
combined multi-frequency IVUS and IVPA is a promising solution to meet the need of
identifying more vulnerability information in both anatomic structure and chemical
composition features. The laser and ultrasonic transducer are two key factors that affect
the IVPA imaging performance. A 500Hz optical parametric oscillator (OPO) laser at
1725nm has been developed specifically for lipid detection by the collaborator group at
52
UCI. The ultrasonic transducer is the critical component of the IVUS-IVPA imaging
catheter since it is responsible for picking up both the photoacoustic and ultrasonic signal.
A recently developed lead indium niobate-lead magnesium niobate-lead titanate (PIN-
PMN-PT) 1-3 Composite material is an extremely viable candidate to renovate the
multimodal imaging catheter for IVUS-IVPA imaging. The ultimate goal is to develop a
fully integrated multi-frequency IVUS-IVPA imaging catheter and system to provide
detailed structure information (by ultrahigh-frequency transducer), lipid content
information (by IVPA at 1.7𝜇𝑚) on a complete picture of the vessel wall (by low-
frequency ultrasound transducer). This chapter describes the progress I have made in the
design, fabrication, and test of the imaging catheter to advance the integrated multi-
frequency IVUS and IVPA imaging technique.
3.2 Method
3.2.1 Probe design and fabrication
For the ultrasonic transducer, the piezoelectric material is the critical component
that will affect imaging performance of the transducer in terms of both resolution and
sensitivity. Electromechanical coupling coefficient (𝑘
!
), piezoelectric strain constant (𝑑
!!
)
and dielectric permittivity (𝜀
!
/𝜀
!
) are three important parameters of the piezoelectric
material. Higher 𝑘
!
and 𝑑
!!
values represent higher energy conversion efficiency of the
material and thus result in higher sensitivity of the transducer. 𝜀
!
/𝜀
!
, together with
aperture size and thickness of the material (determined by the center frequency),
53
determines the electrical impedance of the transducer. Considering the small aperture size
(400~500 𝜇𝑚 ) required for the IVUS application, materials with high dielectric
permittivity (𝜀
!
/𝜀
!
) are preferred to achieve 50Ω electrical impedance matching between
imaging probe and imaging system electronics. The single crystal Pb(Mg
1/3
Nb
2/3
)-PbTiO
3
(PMN-PT) (H.C. materials, Bolingbrook, IL) with high 𝑘
!
(0.58), 𝑑
!!
(2000 pCN
-1
) and
𝜀
!
/𝜀
!
(5229) is a good candidate to fabricate miniature acoustic element, which makes it
favorable for IVUS transducer (Zhou et al., 2007). The recently developed PIN-PMN-PT
single crystal 1-3 composite material (Figure 3-4) shows promising potential to renovate
the ultrasound transducer owing to its superior properties (X. Li et al., 2014). First, it has
higher electromechanical coupling coefficient, which can increase the bandwidth and
then improve the image resolution. Second, it has lower acoustic impedance, which
improves impedance matching and increases sensitivity. Third, it has higher thermal
stability, which is especially important for IVPA imaging because of the heating effect
caused by the laser pulse. I evaluated the performance of this composite material by
comparing it with the commonly used PMN-PT for intravascular photoacoustic imaging.

Figure 3-4: Structure of 1-3 composite material. (adpated from http://www.neptune-
sonar.co.uk/)

54
In my current study, both PIN-PMN-PT single crystal 1-3 composite and PMN-
PT were used to fabricate the imaging probes for performance comparison in IVPA
imaging. The PMN-PT acoustic stacks (Figure 3-5) at 45MHz and 100MHz were
prepared following the general fabrication procedures (Cannata, Ritter, Chen, Silverman,
& Shung, 2003a) and then diced along thickness direction into 0.4x0.4mm square shape
acoustic stack. The 45MHz PIN-PMN-PT 1-3 composite acoustic stacks were customized
by company (HC materials, Bolingbrook, IL, USA). A 50Ω coaxial cable of 250𝜇𝑚 outer
diameter was used to electronically bridge the acoustic stack to other electronics with
core wires and shielding wires connected to the conductive backing layer and matching
layer, respectively. Finally, a thin parylene layer is deposited over the entire probe for
additional acoustic matching and electronic protection.

Figure 3-5: Layer structure of the acoustic stack for the imaging catheter.

Probe design is also an important factor for multimodal intravascular imaging. I
need to consider both the size limitation (~1mm) and image quality requirement. Because
of the size constraints, the same ultrasound transducer is used for picking up both
ultrasonic and photoacoustic signal. Also, the ultrasonic transducer and optical fiber were
55
fixed sequentially in a thin-wall stainless tubing of 1mm outer diameter (OD) with a
window for the optical and acoustic beam to travel, as illustrated in Figure 3-6. The light
from 1725nm OPO laser was coupled into a 105𝜇𝑚 core multimode fiber (MMF). The
fiber tip was polished to have a 38° angle and then housed in a quartz cap for a fiber-air
interface to ensure total internal reflection within MMF at the fiber tip. Type II probe
design is the modified version of Type I probe aimed at improving the acoustic and
optical beam overlap region by decreasing the distance offset between ultrasonic
transducer element and optical fiber tip. The transducer was tilted towards the fiber cap to
maximize the optical/acoustic beam overlap in the field of view. The coaxial cable and
MMF were then enclosed in a double-wounded flexible torque coil and connected to the
slip ring and optical rotary joint, respectively. The torque coil shaft translates the
rotational torque from the rotary motor to the probe tip during imaging. The total optical
insertion loss from laser-fiber coupling and rotatory joint coupling is around -3dB. The
output power at the catheter tip is around 60mW, which corresponds to 120 µμJ/pulse at
500Hz pulse repetition frequency (PRF).

56

Figure 3-6: Photography and schematics of integrated IVUS-IVPA probe of two different
types of design.
3.2.2 Imaging System Setup
The schematic diagram of the integrated IVUS-IVPA imaging system setup is
illustrated in Figure 3-7. The custom-built OPO laser emits light to the catheter tip and in
the meantime, sends a master trigger signal 𝑇
!
to the delay generator. The delay generator
sends trigger 𝑇
!
to the 12-bit digitizer with sampling rate as high as 3GS/s (Gage Applied
Technologies, Lockport, IL, USA). The photoacoustic signal is digitized before the
pulser/receiver is triggered. Synchronized by delay generator, the pulser/receiver is
activated 14𝜇𝑠 later to acquire ultrasound signal for the same A-line after photoacoustic
signal. The time needed for each A-line with photoacoustic and ultrasound signal
corresponding to 6mm depth is 22𝜇𝑠 at a 500MHz sampling rate. A rotary joint device is
built for motion control and signal coupling between rotational and stationary
components. The device is composed of a rotational motor (Animatics, Santa Clara, CA,
USA), an optical rotary joint (Princetel, Inc., Pennington, New Jersey) and a brushed slip
ring (Prosperous, Co., Hangzhou, China). The optical rotary joint and brushed slip ring
are for optical and electrical signal coupling, respectively. All these parts are mounted on
57
a linear stage for the pullback motion during 3D imaging. With a 500Hz pulse repetition
frequency and one revolution per minute (RPM) rotation speed, 2D image data with 500
A-lines are acquired at one frame per second. A custom-developed LabView program
(National Instruments, Austin, TX) was used to control the imaging procedure and save
the raw image data for offline processing and image display.

Figure 3-7: Schematic diagram of the IVUS-IVPA imaging system.
3.3 Experimental Imaging Results
In order to evaluate the imaging performance of the developed imaging catheter, I
conducted imaging experiments through both in vitro phantom imaging and ex vivo
human artery imaging. The imaging results demonstrate the promising potential of this
integrated IVUS-IVPA imaging technology for comprehensive characterization of
atherosclerotic plaque.
58
3.3.1 Lipid-Mimicking Phantom Imaging
A cylindrical agar-based tissue-mimicking phantom with a 3.5mm diameter
lumen in the center was fabricated. silicon dioxide powder was added to mimic the
acoustic scattering effect of tissue. A lumen of 2mm diameter filled with butter was
located at 3mm away from the center, acting as a lipid-mimicking phantom to evaluate
the lipid-selective IVPA imaging using the laser at 1725nm. The phantoms were imaged
in the deionized water bath using both 45MHz PMN-PT probe (Type I and Type II) and
45MHz PIN-PMN-PT single crystal 1-3 composite probe (Type I). The imaging results
are displayed in Figure 3-8. In the IVUS image, the butter inclusion can be detected as
the low echogenicity area with moderate contrast. In the IVPA image, only the butter
inclusion was detected, which demonstrates the high image contrast of IVPA imaging
with a 1725nm laser for lipid detection. With the same system setup and probe design,
the composite probe provides slightly higher signal-to-noise ratio (SNR) than PMN-PT
probe (12dB vs 10dB). The Type II PMN-PT imaging probe was also evaluated to
investigate the performance of modified probe design. The stronger photoacoustic signal
can be distinctively observed in Figure 3-8(g) compared with Figure 3-8(a). The SNR
was increased to 18dB with Type II probe design.

59

Figure 3-8: Lipid-mimicking phantom imaging. (a) IVUS, (b) IVPA, and (c) combined
IVUS-IVPA images of the phantom by 45MHz PMN-PT probe of Type I; (d) IVUS, (e)
IVPA, and (f) combined IVUS-IVPA images of the phantom by 1-3 composite probe of
Type I; (g) IVUS, (h) IVPA, and (i) combined IVUS-IVPA images of the phantom by
45MHz PMN-PT probe of Type II.
3.3.2 Ex vivo Human Artery Imaging
The IVPA imaging capability of the probes was also evaluated by imaging the
human carotid artery. Type I probe failed to detect the lipid signal from the artery
because of lower IVPA sensitivity. Type II probe was able to detect the lipid from the
60
atherosclerotic plaque and the acquired images are displayed in Figure 3-9. The intima
thickening is clearly delineated in IVUS image (Figure 3-9(b)) at 3 o’clock direction but
the lipid content cannot be identified clearly as indicated by the red arrowheads. The
photoacoustic signal from lipid constituents at the corresponding location is shown in
IVPA image (Figure 3-9(a)) as indicated by the white arrowheads. By overlaying the
IVPA image onto IVUS image, the co-registered image (Figure 3-9(c)) provides both
structural and compositional information necessary for vulnerable plaque detection.

Figure 3-9: Ex vivo human cadaver carotid artery imaging. (a) IVPA image, (b) IVUS
image, and (c) combined IVUS-IVPA image of human carotid artery using 45MHz
PMN-PT Type II probe.

To validate the IVUS imaging performance of the 45MHz and 100MHz
ultrasound transducers, human cadaver coronary artery was imaged in the deionized
water bath and the acquired images are shown in Figure 3-10. The 45MHz transducer
provides complete anatomical information of the coronary artery wall owing to the deep
penetration depth, as illustrated in Figure 3-10(a). As shown in Figure 3-10(b), although
the 100MHz transducer can only penetrate through the intima layer due to the limited
penetration depth, the image acquired by the 100MHz transducer has both improved axial
61
and lateral resolution. The fused image pair acquired by 45-100MHz imaging catheter is
shown in Figure 3-10(c). The result suggests that the multi-frequency IVUS imaging
catheter can achieve a comprehensive visualization of the artery wall by integrating the
deep penetration depth of low-frequency transducer and high resolution of high-
frequency transducer.

Figure 3-10: Ex vivo human cadaver coronary artery imaging. (a) 45MHz, (b) 100MHz
and (c) combined multi-frequency IVUS image of human coronary artery.
3.4 Conclusions and Discussions
The preliminary results from both in vitro and ex vivo imaging present the
promising potential of the integrated multi-frequency IVUS-IVPA imaging. However,
several aspects need to be improved to develop the fully integrated probe with better
image quality. First, the optical pulse energy in the current system (120µμJ) is not high
enough to activate strong photoacoustic signal. The laser system should be fine-tuned to
get lower coupling loss and higher optical output. Second, the current imaging speed
(1Hz) is still insufficient for real-time imaging required by the clinical application (>
20Hz). The imaging speed is limited by the laser. A 1725 nm laser with 10kHz PRF is
62
under development to achieve higher imaging speed in the near future. Third, the design
of the imaging probe needs to be modified by focusing the optical beam using gradient-
index (GRIN) lens to further increase the optical intensity. The schematic diagram of the
focused probe design is shown in Figure 3-11(a). Fourth, in terms of experiment design,
lipid-mimicking material with photoacoustic response closer to the plaque lipid should be
adopted to better evaluate the imaging performance of the imaging catheter.
Ultimately, the fully integrated imaging probe with two ultrasound transducer
elements and one optical fiber will be designed and evaluated. The schematic diagram of
a possible multi-frequency IVUS-IVPA probe design is depicted in Figure 3-11(b). The
two transducer elements are aligned back-to-back, which not only ensures the co-
registered IVUS images and but also fits into the size constraints. The focused optical
beam from the GRIN lens is reflected by the angled rod mirror to overlap the ultrasonic
beam from the low-frequency transducer element. Three frames of image data are
acquired during each revolution of the integrated probe: the low-frequency IVUS image
provides the complete picture of the vessel wall which helps to diagnose the plaque
burden; the high-frequency IVUS image is capable of differentiating the thin fibrous cap
over the lipid necrotic core to stratify the vulnerability of the atherosclerotic plaque; the
lipid-specific IVPA image maps the lipid distribution under thin fibrous cap onto the
IVUS image. With only one imaging probe through percutaneous coronary intervention
(PCI) procedure, the two important diagnostic makers of the typical vulnerable plaque
TCFA can be reliably detected, which will better inform interventional cardiologists for
both diagnosis and the following treatment.
63

Figure 3-11: Schematic diagram of future probe design: (a) Focused IVUS-IVPA probe;
(b) Fully integrated multi-frequency IVUS-IVPA probe.

64
Chapter 4 Intravascular Ultrasound-based Single Beam
Acoustic Tweezer
4.1 Introduction
Intravascular ultrasound (IVUS) is a well-known catheter-based technology for
diagnostic intravascular imaging. It helps to visualize the whole artery wall structure,
guide percutaneous coronary intervention (PCI) and assess treatment outcome for
cardiovascular disease. More recently, the combined use of IVUS with microbubbles has
been applied to perform intravascular molecular imaging and enhance drug, gene and
stem cell delivery efficiency (Phillips, Klibanov, Wamhoff, & Hossack, 2012b) (Kilroy et
al., 2015) (Hossack, 2015) (J. Kim et al., 2017). A schematic diagram of therapeutic
IVUS catheter using sirolimus-loaded microbubbles is shown in Figure 4-1. To overcome
the challenges posed by the blood flow, especially in large vessels, the primary acoustic
radiation force was utilized to enhance the concentration of microbubbles near the vessel
wall (J. P. Kilroy, A. L. Klibanov, B. R. Wamhoff, & J. A. Hossack, 2012). The primary
acoustic radiation force, resulting from the momentum transfer between acoustic wave
and particles in the acoustic field, helps to accumulate the ligand-labeled and therapeutic
agent-laden microbubbles by moving them from the blood stream onto the endothelial
layer of a vessel wall. Both microbubble-based molecular imaging and therapeutic
delivery can be improved through this mechanism owing to the enhanced binding
between microbubbles and targeted site (Phillips, Dhanaliwala, Klibanov, Hossack, &
Wamhoff, 2011) (Rychak, Klibanov, Ley, & Hossack, 2007) (Kilroy, Klibanov,
65
Wamhoff, Bowles, & Hossack, 2014). However, simply moving the microbubbles away
from IVUS catheter along beam axis cannot ensure manipulation in the transverse
direction or along the vessel wall to precisely target the diseased location. A tool that
allows trapping and manipulation of micro-objects in a controllable manner along blood
vessel wall is able to further enhance adhesion of functionalized microbubbles, which is
of significant importance for enhancing molecular imaging and therapy delivery (Meng et
al., 2011) (Lo, Kang, & Yeh, 2015).

Figure 4-1: Schematic diagram of therapeutic IVUS catheter with sirolimus-loaded
microbubbles. The microbubbles are injected through the imaging catheter and then
displaced onto vessel wall by radiation force from IVUS transducer. (adapted from
Kilroy et al., 2015)

Multiple manipulation methods using acoustic wave have been proposed to trap
and position microscopic objects in a controllable manner (Lo et al., 2015) (Shi et al.,
2009)(Choe, Kim, Shung, & Kim, 2011) (Zhu et al., 2016). Yeh et al. experimentally
demonstrated the transverse trapping and manipulation capability of acoustic vortex
produced by a four-element transducer (Figure 4-2). The adjacent planar elements of the
four-element transducer was driven under a pi/2 phase difference along the beam axis and
thus produces the destructive interference pattern that forms a potential well. The high
66
pressure gradient of the potential well confines the microbubbles within the beam and
realizes spatiotemporal manipulation. The best trapping performance is expected in front
of one-fourth of Rayleigh distance (RD) owing to higher transverse force (Lo et al., 2015).
But the potential of this technique for intravascular application will be hindered by the
difficulty in miniaturizing the multi-element transducer. Recently, a three-dimensional
acoustic tweezer based on 3D trapping node created by standing surface acoustic wave
(SSAW) has been developed (Figure 4-3), which can be potentially applied for
bioprinting (Guo et al., 2016). Two orthogonal pairs of interdigital transducers (IDTs) are
used to produce orthogonally superimposed pairs of SSAWs within a microfluidic
chamber. The 3D distributed acoustic field and induced acoustic streaming create an
array of 3D trapping nodes that trap microscale objects. Transverse and vertical
manipulation can be realized by tuning the phase angle of SSAW and the input acoustic
power, respectively. However, the intravascular application of this technique is still
limited because of the bulkiness of the system setup.

Figure 4-2: Trapping model of acoustic vortex: the four-element transducer creates
potential-well that provides trapping force to manipulate microbubbles. (adapted from
(Lo et al., 2015))

67

Figure 4-3: Experimental setup (a) and acoustic trapping node (b) of surface acoustic
wave based 3D manipulation trapping technique. (adapted from Guo et al., 2016)

As the acoustic analog of optical tweezer, single beam acoustic tweezer (SBAT)
can trap and move the micro-objects using the transverse acoustic radiation force
produced from momentum transfer between a tightly focused acoustic beam and micro-
objects (Figure 4-4). Since the SBAT was first demonstrated by Lee and Shung (Lee, Ha,
& Shung, 2005) (Lee & Shung, 2006), its trapping capability has been demonstrated
using single element transducer and phase array (Hsu et al., 2012) (Lam et al., 2013)
(Zheng et al., 2012). In comparison with other acoustic manipulation methods, focused
high-frequency transducer based SBAT provides stronger trapping force and deeper
penetration depth. Therefore it is a good candidate as a safe, cost-effective technique for
intravascular trapping and manipulation. Moreover, in the intravascular imaging setting,
the performance of SBAT will not be compromised by the attenuation or distortion of
tissues over the targeted blood vessel. Li et al. reported the feasibility of in vivo
application of SBAT by demonstrating trapping capability of SBAT after penetrating
rabbit aorta and on a curved surface, respectively, by linearly translating the transducer
(Y. Li, Lee, Chen, Zhou, & Shung, 2014). Although the trapping experiments conducted
68
in this study confirm the potential of SBAT for in vivo application, they are not sufficient
to demonstrate the trapping capability along blood vessel wall in IVUS imaging setting,
where the transducer is rotated for cross-sectional imaging.

Figure 4-4: Schematic diagram of single beam acoustic trapping.

Inspired by this unmet need, a 50MHz focused single element transducer was
designed, fabricated and characterized. The feasibility of intravascular trapping and
manipulation of microscopic objects by rotationally manipulating the IVUS-based SBAT
was initially demonstrated within a mylar film tube filled with distilled water. To
evaluate the biological effects of this intravascular SBAT, mechanical and thermal index
were also estimated based on hydrophone tests. Both mechanical and thermal index meet
the standard of the commercial diagnostic ultrasound system, which further proves its
capability for in vivo intravascular application. This intravascular manipulation method
could be integrated into an IVUS catheter for more efficient therapy delivery to specific
atherosclerotic plaque area.
69
4.2 Method
A 50MHz lithium niobate (LiNbO
!
) side-looking transducer was designed and
fabricated. The transducer was press-focused(Cannata, Ritter, Chen, Silverman, & Shung,
2003b) at a focal length of 3.5mm to achieve an f-number (F
#
) of 1.16. To characterize
the transducer, a pulse-echo test was conducted in deionized (DI) water with an X-cut
quartz plate as a reflecting target. The quartz plate was located at the focal point of the
press-focused transducer. The frequency characteristics were analyzed based on the
normalized frequency spectrum from the pulse-echo test. The beam profile and acoustic
output power of the transducer were measured by a calibrated hydrophone following the
measurement protocol in ‘Acoustic Output Measurement Standard for Diagnostic
Ultrasound Equipment, Revision 3 (NEMA Standards Publication UD 2-2004)’.

Figure 4-5: Experimental configuration for demonstrating intravascular trapping
capability of SBAT.

70
The experimental setup for demonstrating the feasibility of intravascular single
beam acoustic trapping with a focused IVUS transducer is shown in Figure 4-5. The
transducer was fixed onto a connector to follow the rotation of the motor (SM17205D,
Moog Animatics, Milpitas, CA, USA). In order to demonstrate the trapping capability
along the curved surface of mylar film tube, the transducer was rotated by the motor at a
speed of 0.05 rotation per minute (0.3 deg/s). The motor was mounted on a 3D linear
manual positioner (OptoSigma, Santa Ana, CA, USA) for position control along X, Y, Z
direction as shown in Figure 4-5. The aperture surface of the transducer was immersed in
deionized water within the transparent mylar film tube. The radius of mylar film tube was
3.5mm to match the focal depth of the transducer. A transparent chamber was used to
hold the mylar film tube and DI water. Polystyrene microparticles (Megabead, NIST
traceable particle size standard, Polyscience, Inc., Warrington, PA) of 10µμm mean
diameter were injected into the tube as trapping targets. Before the acoustic trapping
experiment, the transducer was focused on the bottom of the mylar film membrane
guided by the pulse-echo test. Then the transducer was moved along the beam axis away
from the bottom by the radius of the polystyrene particles. In this way, the center of
particles was located at the geometrical focus the ultrasound beam. During the trapping
experiment, the transducer was excited by a 25Vpp 50MHz sinusoidal burst from a
function generator (AFG3251, Tektronix, Anaheim, CA, USA) followed by a 50dB RF
power amplifier (525LA, ENI, Rochester, MN). The pulse repetition frequency (PRF)
and duty factor are 1kHz and 1%, respectively. A mercury lamp delivered light to the
polystyrene microparticles through the objective. The motion of microparticles resulting
from the acoustic tweezer manipulation along the inner surface of mylar film tube could
71
be observed through an inverted microscope (IX-71, Olympus, Japan) with 10X objective
and recorded by the CMOS camera (ORCA-Flash2.8, Hamamatsu, Japan). The images
and videos were acquired at the frame rate of 10 frames/s and then saved to a computer.
4.3 Experimental Results
4.3.1 Transducer Characterization

Figure 4-6: Time-domain pulse-echo signal (black solid line) and normalized frequency
spectrum (red dashed line) of the 50MHz transducer.

Figure 4-6 shows the pulse-echo measurement result of the transducer. The
fractional bandwidth of the transducer is 77.2% with a center frequency at 50.5 MHz,
covering a frequency band from 31MHz to 70MHz. 2D beam profile and contour plot of
the transducer measured through hydrophone test are shown in Figure 4-7 with position
(0,0)µm referring to the focal point. The -3dB beam width along the lateral and axial
direction is 50µm and 410µm, respectively. To evaluate the bioeffects of SBAT for
72
intravascular application, derated Spatial Peak Temporal Average Intensity (𝐼
!"#$.!
),
derated Spatial Peak Pulse Average Intensity (𝐼
!""#.!
), Mechanical Index (MI), and
Thermal Index (TI) at the focal point under the excitation conditions during the trapping
experiment were calculated and summarized in Table 4-1(Y. Li et al., 2014)

(Acevedo &
Das-Gupta, 2002)

(Bigelow et al., 2011). All four parameters meet the biosafety standard,
demonstrating that this IVUS-based SBAT is suitable for in vivo intravascular
application.

Figure 4-7: (a) The two-dimensional beam profile on the XZ plane with (0,0) position
referring to the focal point of the transducer. (b) The contour plots delineate pressure
level of -3dB, -6dB, -9dB and -12 dB.

Table 4-1: Acoustic Output Exposure Level of SBAT at focus

I
!"#$.!

(mW/cm
!
)
I
!""#.!

(W/cm
!
)
Mechanical
Index
Thermal
Index
SBAT 386 37 0.13 0.03
Biosafety Standard 430 190 1.9 2

(a) (b)
73
4.3.2 Force Analysis during Manipulation
Disregarding the friction force and the stokes drag force (induced by acoustic
streaming), the microparticles on the tube wall mainly experience 4 forces: acoustic
radiation force (ARF), gravitational force (𝐹
!
), supporting force (𝐹
!
) from mylar film
tube, and buoyancy force (𝐹
!
), as illustrated in Figure 4-8. For particles on the tube wall,
the net force along the ultrasound beam axis is zero, balanced by axial acoustic radiation
force, supporting force, axial component of buoyancy force and gravitational force.
Transverse acoustic radiation force (𝐴𝑅𝐹
!
) varies according to the relative position of
microparticles in ultrasound beam and it always acts in the direction to pull the objects
towards the beam axis within a certain distance range (Zhu et al., 2016)(Azarpeyvand &
Azarpeyvand, 2013). Along the transverse direction, interaction among transverse
acoustic radiation force, transverse components of buoyancy force and gravitational force
enable the manipulation of microparticles along the curved surface. During the
manipulation process, there are three main types of motion: acceleration, constant speed,
and deceleration (Y. Li et al., 2014). The acceleration period is critical because the SBAT
needs to produce enough force to keep the microparticles in the trap. In the initial
acceleration period of manipulation, the forces along transverse direction follow
Newton’s second law and can be described by:
𝐴𝑅𝐹
!
− 𝐹
!
−𝐹
!
sinθ= ma 4-1
where m and a are the mass and acceleration of microparticles, respectively. As shown in
Figure 4-8, to achieve 360° manipulation of microparticles along the vessel wall, the
trapping force has to be adßjusted to be greater than F
!
−F
!
+ma, where θ is 90°.
74

Figure 4-8: Illustration of forces exerting on a microparticle (red circle) on tube wall. The
black dashed line denotes the ultrasound beam axis. Gold dashed circle in the tube center
indicates the rotation trajectory of transducer aperture.
4.3.3 Trapping and Manipulation Results
The focused transducer was demonstrated as capable of trapping the 10µμm
microparticles along the circular surface that is geometrically similar to the blood vessel
wall. The trapping results of a single 10µμm microparticle and a group of 10µμm
microparticles are illustrated in Figure 4-9 and Figure 4-10, respectively. In Figure 4-9,
the bright area in the bottom is the projection of the mylar film tube and the bright
circular pattern is the projection of the aperture surface of the transducer. The green
dashed ellipse is the slot at the bottom of the chamber and can be used as the stationary
reference. It has been demonstrated that the trapped single microparticle within the red
dashed circle can follow the motion of the transducer to zigzag along the circular surface.
In Figure 4-10, the bright area in the top is the projection of the mylar film tube. The
trapped microparticles marked by the red dashed circle cannot only move along the tube
75
axis but also follow the rotation of the motor to either climb up or slide down along the
curved surface of the mylar film tube. The aggregated particles indicated by the green
dashed circles were beyond the trapping range of the SBAT and thus stay unaffected.
These experimental results suggest that IVUS transducer based SBAT can perform
microscopic objects manipulation along the curved surface.

Figure 4-9: Single polystyrene microparticle of 10µμm mean diameter was trapped and
manipulated by the SBAT along mylar film tube. The red dashed circle marks the trapped
microparticle. The green dashed ellipse indicates a reference position to show the motion
of the microparticle. The blue arrow denotes the moving direction of the microparticle.

(a) (b)
(f) (e) (d)
(c)
X Y
Z
X
Y
100um 100um 100um
100um 100um 100um
76

Figure 4-10: A group of polystyrene microparticles of 10µμm mean diameter were trapped
and manipulated by the SBAT along the mylar film tube. The red dashed circle marks the
trapped microparticles. The green dashed circle indicates a reference point to show the
motion of the microparticles. The blue arrow denotes the moving direction of the
aggregated microparticles.
4.4 Conclusions and Discussions
The experimental results reported demonstrate the trapping capability of IVUS-
based SBAT along curved surface under rotational motion. Grounded in these promising
results, much work needs to be done to modify this IVUS-based SBAT for practical
applications. First, in the current therapeutic IVUS technology, microbubbles instead of
solid microparticles are used to load therapeutic agents. Therefore, the trapping and
manipulation of microbubbles need to be demonstrated to further validate the
effectiveness of this technique. Second, the SBAT needs to overcome the viscous drag
force exerted by the blood flow for in vivo application. The effect of viscous drag force
from the blood flow will be investigated in the near future. In summary, I have first
demonstrated the single beam acoustic trapping along a curved surface with an IVUS
X
Y
X Y
Z
(a) (b) (c)
(d) (e) (f)
77
probe, which shows great potential for intravascular contactless manipulation. The
intravascular SBAT shows potential for trapping and manipulating functionalized micro-
objects (labeled with targeted ligand and therapeutic agents). By moving the therapeutic
micro-objects to the plaque area detected by IVUS imaging, the IVUS-based SBAT will
significantly improve efficiency and effectiveness of therapy delivery.

78
Chapter 5 Conclusions and Future Work
5.1 Conclusions
In summary, this dissertation has reported the efforts in improving intravascular
ultrasound (IVUS) based technology by improving image quality, developing an
integrated multimodal imaging technique, and exploring the capability to enhance
therapeutic delivery.
The combined use of virtual source synthetic aperture (VSSA) and coherence
factor weighting (CFW) is applied to enhance the image quality of the single element
IVUS transducer over the entire field of view, which cannot be achieved using previous
methods including increasing the frequency and mechanical focusing. The proposed
method VSSA-CFW was evaluated in both Field II simulation and experiments through
imaging of wire phantom, homogeneous agar phantom, anechoic cyst phantom, and
human cadaver artery. The results suggest that VSSA-CFW can effectively improve
lateral resolution and contrast-to-noise ratio (CNR) of the IVUS image in both the cross-
sectional view and the longitudinal view. Additionally, the influence of non-uniform
rotational distortion (NURD) on the performance of VSSA-CFW has been evaluated
through the Field II simulation. VSSA-CFW is capable of improving CNR with moderate
NURD. Moreover, VSSA-CFW can be readily integrated into current IVUS imaging
systems by adding the post-processing module after RF data acquisition.
To get more diagnostic information of the vulnerable plaque (thin-cap
fiberoatheroma, TCFA), multi-frequency IVUS and lipid-selective intravascular
79
photoacoustic (IVPA) are integrated to provide a deep penetration image of entire vessel
wall using low-frequency IVUS (45MHz), a high-resolution image of the intima layer
using high frequency IVUS (100MHz), and a plaque lipid selective image using IVPA at
1.7𝜇𝑚 spectral band. The integrated imaging catheter is designed, fabricated and
evaluated through both in vitro phantom imaging and ex vivo human artery imaging. The
images acquired from the imaging catheter demonstrated the capability of high resolution,
deep penetration, and lipid specificity.
To enhance the therapeutic delivery efficiency, the manipulation capability of the
IVUS-based single beam acoustic tweezer (SBAT) was investigated. A 50MHz side-
looking transducer was designed, fabricated and evaluated. The microparticles can be
manipulated by the designed transducer along the curved surface of a mylar film tube
wall. The acoustic output during the trapping experiment was also measured and all the
parameters (I
!"#$.!
, I
!""#.!
, Mechanical Index, Thermal Index) meet the safety standard.
Therefore, the IVUS-based SBAT holds the potential to manipulate microscopic targets
along blood vessel wall for higher therapeutic uptake.
5.2 Future Work
The end goal of my study is to advance the IVUS technology to a robust
theranostic tool that can provide high image quality, comprehensive diagnostic
information (anatomic and compositional), and efficient therapy delivery within a single
intervention procedure. Though the study results presented in this dissertation
demonstrate the promising potential to advance current IVUS technology, future
80
development is required to further enhance the proposed methods and bring in clinical
value. For the VSSA-CFW method, its performance on in vivo IVUS image data acquired
during rotational pullback motion still needs to be studied to validate its effectiveness
under motion artifact and NURD. For the multi-frequency IVUS-IVPA imaging, a 1.7um
laser with high speed and high output energy needs to be developed to improve the
imaging speed and sensitivity. Then, a fully integrated and focused imaging catheter
should be prototyped to comprehensively evaluate the image quality of this multimodal
technique. For the IVUS-based SBAT, the work presented in Chapter 4 is a preliminary
demonstration of its principle and concept. Further investigation on the trapping
capability of functionalized microbubbles within blood flow is necessary to confirm the
feasibility of implementing this technique for clinical application.

81
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Asset Metadata

Creator Yu, Mingyue (author)

Core Title Advamces in intravascular ultrasound (IVUS)-based technology

Contributor Electronically uploaded by the author (provenance)

School Andrew and Erna Viterbi School of Engineering

Degree Doctor of Philosophy

Degree Program Biomedical Engineering

Publication Date 01/20/2018

Defense Date 11/20/2017

Publisher University of Southern California (original), University of Southern California. Libraries (digital)

Tag imaging,intravascular ultrasound (IVUS),OAI-PMH Harvest,photoacoustic imaging,virtual source synthetic aperture

Language English

Advisor Zhou, Qifa (committee chair), Chen, Mike Shuo-Wei (committee member), Shung, K. Kirk (committee member)

Creator Email mingyue930@gmail.com,mingyuey@usc.edu

Permanent Link (DOI) https://doi.org/10.25549/usctheses-c40-464134

Unique identifier UC11268244

Identifier etd-YuMingyue-5970.pdf (filename),usctheses-c40-464134 (legacy record id)

Legacy Identifier etd-YuMingyue-5970.pdf

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Document Type Dissertation

Rights Yu, Mingyue

Type texts

Source University of Southern California (contributing entity), University of Southern California Dissertations and Theses (collection)

Access Conditions The author retains rights to his/her dissertation, thesis or other graduate work according to U.S. copyright law. Electronic access is being provided by the USC Libraries in agreement with the a...

Repository Name University of Southern California Digital Library

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Abstract (if available)

Abstract Cardiovascular disease (CVD), the number one killer in the United States, causes around 30.8% of all deaths. Most of CVD is related to a progressive process, atherosclerosis. Atherosclerotic plaque builds up gradually within the artery wall and may remain asymptomatic until high-risk plaque (vulnerable plaque) ruptures, resulting in heart attack or even sudden death. Intravascular ultrasound (IVUS) has frequently been used as a clinical diagnostic tool for intravascular imaging to measure artery lumen dimension and the total plaque volume. However, current IVUS technique is not sufficient for accurate and robust diagnosis of vulnerable plaque, especially regarding image quality and image information. In addition to the diagnostic benefit, the therapeutic value of IVUS in the combined use of microbubbles has also been intensively investigated in recent years. However, a technique for controllable trapping and manipulation of microscopic targets along artery wall is still needed for efficient therapy delivery. ❧ In this dissertation study, IVUS technique has been advanced in three aspects: 1. applied virtual source synthetic aperture (VSSA) focusing and coherence factor weighting (CFW) to improve IVUS image quality