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High-frequency ultrasonic transducers for photoacoustic applications

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High-frequency ultrasonic transducers for photoacoustic applications

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HIGH-FREQUENCY ULTRASONIC TRANSDUCERS FOR PHOTOACOUSTIC
APPLICATIONS

by

Ruimin Chen

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 2014

Copyright 2014 Ruimin Chen

ii

Dedication

To my beloved

Parents
Jian Chen & Liming Yu

iii

Acknowledgements
I would like to thank the National Institute of Health (NIH) for its generous
funding in support of the work presented in this dissertation. This four years journey
through graduate school has brought me in contact with many wonderful people who
have contributed immensely to my personal development as a scientist and engineer, and
to the development of ultrasonic transducers for photoacoustic applications.
First, I am sincerely grateful to my committee members. Foremost, my advisors,
Dr. K. Kirk Shung and Dr. Qifa Zhou, their guidance, mentoring and support since the
day I entered into the laboratory has allowed me to excel and be passionate about this
project. They are truly great mentors who pass to me not only their knowledge but also
their life philosophy. I would also like to express my sincere gratitude to my Dr. Jesse
Yen and Dr. Yong Chen for their suggestion during my preparation of this dissertation.
I am greatly thankful to all my colleagues and friends at the Resource Center for
Medical Ultrasonic Transducer Technology, especially current and graduated students
Dr. Dawei Wu, Dr. Fan Zheng, Dr. Xiang Li, Dr. Hsiusheng Hsu, Mr. Teng Ma and Mr.
Nestor Cabrera, our lab manager Dr. Hyung-Ham Kim our previous lab manager Dr.
Jonathan Cannata, transducer engineer Jay Williams, previous postdoctoral fellows Dr.
Sien Ting Lau and Dr. Kwok Ho Lam, and our budget analyst Peter Lee. We have been
working together for many years and shared a lot pains and happiness. They supported
me throughout my study and made my stay at the center an enjoyable experience.
My work would not have been realized without the help of Dr. Changgeng Liu
from Geospace Research, Inc. He provided valuable helps on ultrasonic array transducer
iv

fabrication process. He also contributed to my knowledge of ultrasonic array design and
fabrication through helpful discussions.
I am heartily thankful to Dr. Lihong Wang from Washington University in St.
Louis, who is my project supervisor in collaboration. Without Dr. Wang’s support, my
research work will never be possible. I would also like to thank my colleagues from
Optical Imaging Laboratory at WUSTL especially Dr. Joon-Mo Yang, Dr. Dakang Yao
and Mr. Chi Zhang.
Last but not the least, I would like to thank my parents who raised me up and give
me unconditional love.

v

Table of Contents
Dedication ........................................................................................................................... ii
Acknowledgements ............................................................................................................ iii
List of Tables .................................................................................................................... vii
List of Figures .................................................................................................................. viii
Abstract ............................................................................................................................. xii
Chapter 1 Introduction ........................................................................................................ 1
Chapter 2 Photoacoustic Theory ......................................................................................... 4
2.1 Photoacoustic Effect ......................................................................................... 4
2.2 Photoacoustic Imaging ...................................................................................... 5
2.3 PA Imaging Modalities in Biomedicine ........................................................... 7
Chapter 3 Ultrasonic Transducer ........................................................................................ 9
3.1 Ultrasonic Transducer Fundamentals ............................................................... 9
3.1.1 Piezoelectric Parameters .................................................................... 9
3.1.2 Piezoelectric Materials ..................................................................... 12
3.1.3 Mechanical Matching....................................................................... 13
3.2 Single-element and Array Transducers ........................................................... 16
3.2.1 Single-element transducer ................................................................ 16
3.2.2 Array Transducer ............................................................................. 18
3.3 Axial and Lateral Resolution .......................................................................... 20
3.4 High-frequency Ultrasound Imaging .............................................................. 21
3.5 Issues of Ultrasonic Attenuation at High-frequency ....................................... 22
3.6 Ultrasonic Transducer Modeling and Testing................................................. 23
3.6.1 Krimholtz, Leedom, and Matthaei (KLM) equivalent circuit model23
3.6.2 Transducer Modeling Softwares ...................................................... 25
3.7 Transducer Performance Evaluation ............................................................... 27
3.7.1 Electrical Impedance Test ................................................................ 27
3.7.2 Pulse-echo Test ................................................................................ 27
3.7.3 Insertion Loss Test ........................................................................... 29
3.7.4 Crosstalk Measurement .................................................................... 30
3.7.5 Directivity Measurement ................................................................. 30
Chapter 4 Highly Focused High-frequency Ring Transducer for Photoacoustic Microscopy
Application ......................................................................................................................... 32
vi

4.1 Introduction to PAM ....................................................................................... 32
4.2 PAM Transducer ............................................................................................. 33
4.2.1 Transducer Design ........................................................................... 34
4.2.2 Transducer Fabrication Process ....................................................... 37
4.2.3 Transducer Performance Characterization ....................................... 38
4.3 UV-PAM System and Imaging ....................................................................... 40
4.4 Summary ......................................................................................................... 47
Chapter 5 High-frequency Ultrasonic Transducer for Photoacoustic Endoscopy Application .. 48
5.1 Introduction to PAE ........................................................................................ 48
5.2 First Generation Transducer Design: Lens-focused Ring Transducer ............ 49
5.2.1 Transducer Design and Fabrication ................................................. 49
5.2.2 Transducer Performance Evaluation ................................................ 54
5.2.3 Ultrasound Resolution PAE Probe Implementation ........................ 55
5.2.4 In vivo PAE Imaging ........................................................................ 57
5.3 Second Generation Transducer Design: Self-focused Ring Transducer......... 60
5.3.1 Transducer Design and Fabrication Process .................................... 60
5.3.2 Transducer Performance Evaluation ................................................ 61
5.4 Third Generation Transducer Design: Self-focused Circular Shape Transducer ....... 62
5.4.1 Transducer Design and Fabrication Process .................................... 62
5.4.2 Transducer Performance Evaluation ................................................ 63
5.4.3 Flexible Shaft-based PAE Probe Implementation ........................... 64
5.5 Summary ......................................................................................................... 65
Chapter 6 High-frequency Phased Array for Photoacoustic Endoscopy Application ...... 67
6.1 Introduction ..................................................................................................... 67
6.2 High-frequency PMN-PT Kerfless Phased Array........................................... 70
6.2.1 Design .............................................................................................. 70
6.2.2 Fabrication Process .......................................................................... 72
6.2.3 Performance Evaluation ................................................................... 75
6.2.4 Imaging Test .................................................................................... 80
6.3 High-frequency PZT Kerfed Phased Array .................................................... 85
6.3.1 Design .............................................................................................. 85
6.3.2 Fabrication Process .......................................................................... 86
6.3.3 Performance Evaluation ................................................................... 88
6.4 Summary ......................................................................................................... 89
Chapter 7 Summary and Future Work .............................................................................. 91
7.1 Summary ......................................................................................................... 91
7.2 Suggestions for Future Work .......................................................................... 92
Bibliography ..................................................................................................................... 95

vii

List of Tables
Table 1-1: Comparison of low-frequency US, high-frequency US, OCT, and PA imaging. .. 2
Table 3-1: Properties of important piezoelectric materials. .............................................. 13
Table 4-1: Design parameters and modeling results of PAM transducer. ........................ 34
Table 5-1: Design parameters and modeling results of 1
st
generation PAE transducer. ... 51
Table 6-1: Properties of PMN-PT single crystal............................................................... 71
Table 6-2: Design parameters of the phased array transducer. ......................................... 72
Table 6-3: Comparison of simulated and measured electrical impedance results. ........... 76
Table 6-4: Comparison of simulated and measured pulse-echo response and its FFT spectrum. . 78
Table 6-5: Comparison between measured and modeling results for the array. ............... 79
Table 6-6: Specifications of 32-channel digital imaging system. ..................................... 80
Table 6-7: Design parameters of the kerfed phased array transducer. .............................. 86
Table 6-8: Phased array performance summary. .............................................................. 90

viii

List of Figures
Figure 2-1: Basic principle of the PA effect. ...................................................................... 5
Figure 3-1: A simplified model of an ultrasonic transducer. ............................................ 14
Figure 3-2: Attenuation of variety tissues in the frequency range from 10 to 100 MHz .. 23
Figure 3-3: The KLM model for a single-element transducer. ......................................... 24
Figure 3-4: Set-up of pulse-echo test. ............................................................................... 28
Figure 3-5: Set-up of insertion loss test. ........................................................................... 29
Figure 3-6: Set-up of array crosstalk measurement. ......................................................... 30
Figure 3-7: Set-up of array directivity measurement. ....................................................... 31
Figure 4-1: Set-up of a conventional PAM system. .......................................................... 33
Figure 4-2: Schematic of PAM transducer. ...................................................................... 34
Figure 4-3: Electrical impedance and phase modeling results of PAM transducer by
PiezoCAD. ..................................................................................................... 35
Figure 4-4: Pulse-echo impulse reponse and its FFT spectrum modeling result of PAM
transducer by PiezoCAD. ............................................................................... 36
Figure 4-5: Field II simulation of the acoustic field transmission of PAM transducer with
a contour representation. ................................................................................ 37
Figure 4-6: Photo of a PAM transducer. ........................................................................... 38
Figure 4-7: Measured electrical impedance and phase of PAM transducer. .................... 39
Figure 4-8: Pulse-echo measurement of PAM transducer. ............................................... 39
Figure 4-9: Absorption spectra of DNA, RNA, and protein. ............................................ 41
Figure 4-10: Schematic of the PAM system with a ring-shape transducer....................... 41
Figure 4-11: Resolution of UV-PAM. .............................................................................. 42
Figure 4-12: PAM image of epithelial cells in the ex vivo lip of a mouse. ....................... 43
Figure 4-13: Photoacoustic image and histological micrograph of cell nuclei. ................ 43
ix

Figure 4-14: In vivo en face photoacoustic images of the skin of mouse ears in the form of
maximum amplitude projection (MAP). ........................................................ 44
Figure 4-15: In vivo en face distribution of cell nuclei in the skin of mouse ears. ........... 45
Figure 4-16: Signal and contrast of in vivo photoacoustic images of cell nuclei versus the
optical wavelength. ........................................................................................ 46
Figure 4-17: In vivo photoacoustic imaging of cell nuclei in the skin of mouse ears with
varied laser pulse energy at a wavelength of 250 nm. ................................... 47
Figure 5-1: Schematic diagram of 1
st
generation PAE transducer. ................................... 50
Figure 5-2: Electrical impedance and phase modeling results of 1
st
generation PAE
transducer by PiezoCAD. ............................................................................... 51
Figure 5-3: Pulse-echo impulse reponse and its FFT spectrum modeling result of 1
st

generation PAE transducer by PiezoCAD. .................................................... 52
Figure 5-4: Field II simulation of the acoustic field transmission with a contour
representation of 1
st
generation PAE transducer. ........................................... 53
Figure 5-5: Photos of a finished 1
st
generation PAE transducer. ...................................... 54
Figure 5-6: Measured electrical impedance and phase of the 1
st
generation PAE transducer.
........................................................................................................................ 55
Figure 5-7: Measured pulse-echo impulse reponse and its FFT spectrum of the 1
st

generation PAE transducer............................................................................. 55
Figure 5-8: Schematic, photograph, and field of view of the photoacoustic endoscope.. 56
Figure 5-9: Illustration of simultaneous, multi-wavelength photoacoustic (PA) and
ultrasonic (US) endoscopy. ............................................................................ 58
Figure 5-10: Simultaneous, co-registered PA and US endoscopy of a rabbit esophageal
tract in vivo over an 18 mm diameter and 15 cm long FOV.. ........................ 60
Figure 5-11: Schematic of 2
nd
generation PAE transducer. .............................................. 61
Figure 5-12: Photos of a finished 2
nd
generation PAE transducer. ................................... 61
Figure 5-13: Measured pulse-echo impulse reponse and its FFT spectrum of the 2
nd

generation PAE transducer............................................................................. 62
Figure 5-14: Photos of a finished 3
rd
generation PAE transducer. ................................... 63
Figure 5-15: Measured pulse-echo impulse reponse and its FFT spectrum of the 3
rd

generation PAE transducer............................................................................. 64
x

Figure 5-16: A diagram of the flexible shaft-based PAE. ................................................ 65
Figure 6-1: Schematic of the phased array transducer. ..................................................... 71
Figure 6-2: The array pattern with 32-elements. .............................................................. 73
Figure 6-3: Fabrication flow of PMN-PT kerfless phased array. ..................................... 74
Figure 6-4: Photo of the PMN-PT kerfless phase array prototype. .................................. 74
Figure 6-5: Simulated electrical impedance and phase of a representative kerfless array
element by PZFlex. ........................................................................................ 75
Figure 6-6: Measured electrical impedance and phase of a representative kerfless array
element. .......................................................................................................... 76
Figure 6-7: Simulated pulse-echo response and its FFT spectrum of a representative
kerfless array element. ................................................................................... 77
Figure 6-8: Measured pulse-echo response and its FFT spectrum of a representative
kerfless array element. ................................................................................... 77
Figure 6-9: PZFlex simulated and measured cross-talk of the array. ............................... 78
Figure 6-10: Measured one-way directivity for a single array element of the kerfless array.
........................................................................................................................ 79
Figure 6-11: The arrangement of five tungsten wire targets. ............................................ 81
Figure 6-12: Wire phantom imaging with the 40 MHz PMN-PT kerfless phased array. . 82
Figure 6-13: Lateral and axial line spread functions for the 2
nd
wire of the wire phantom.
........................................................................................................................ 82
Figure 6-14: Field II simulated wire phantom image with 40 MHz phased array. ........... 83
Figure 6-15: Field II simulated lateral and axial line spread functions for the 2nd wire of
the wire phantom. ........................................................................................... 83
Figure 6-16: Co-registered US and PA image of a graphite rod using the 40-MHz PMN-
PT single crystal kerfless phased array. ......................................................... 84
Figure 6-17: PZT-5H 2-2 composite................................................................................. 86
Figure 6-18: Flex-circuit for the kerfed array. .................................................................. 87
Figure 6-19: The front surface of the kerfed array. .......................................................... 87
xi

Figure 6-20: Measured electrical impedance and phase of a representative kerfed array
element. .......................................................................................................... 88
Figure 6-21: Measured pulse-echo response and its FFT spectrum of a representative
kerfled array element. .................................................................................... 89

xii

Abstract
Photoacoustic imaging is a novel hybrid imaging technique that combines the
virtues of both optics and ultrasound by providing the high contrast of optical imaging,
while retaining the high resolution and deep depth imaging capabilities of ultrasonic
imaging. Because of these advantages, recently there is much interest in developing novel
photoacoustic imaging modalities.
Ultrasonic transducer is very critical among all components of photoacoustic
imaging system and can determine photoacoustic image quality. Similar to ultrasound
imaging, the resolution and imaging depth of photoacoustic imaging is scalable, depending
on characteristics of the ultrasonic transducer used, such as frequency, bandwidth and
shape.
This research investigated ultrasonic transducers for two photoacoustic imaging
modalities: photoacoustic endoscopy and photoacoustic microscopy. Different types of
high-frequency single element and array ultrasonic transducers have been designed and
fabricated for each modality. Based on the specially designed ultrasonic transducers,
photoacoustic endoscopy imaging probe and photoacoustic microscopy imaging system
have been built. Ex vivo and in vivo studies have been carried out to validate these new
imaging concepts. The imaging results strongly supported the enhancements brought by
the new transducers and also suggested potential improvements in future.

1

Chapter 1 Introduction
The development of novel biomedical imaging modalities is stimulated by the
manifested need for high speed, high resolution, and non-invasive techniques. Laser
induced photoacoustic (PA) imaging is an imaging modality based on the intrinsic optical
properties of biological tissue and ultrasonic detection at high-frequencies (>20MHz) (Xu
and Wang 2006). PA imaging can be used as a diagnostic tool which combines the strengths
of both optical and acoustic imaging techniques.
PA imaging has its physical basis in a phenomenon called the PA effect. In
biological tissue, incident laser light will experience both scattering and absorption. The
PA effect occurs when the pulsed light energy is absorbed locally in the tissue, causing a
small rapid temperature rise in the medium, and inducing thermoelastic expansion. This
expansion produces pressure transients, which propagate as acoustic waves throughout the
tissue. A PA wave is unique in the sense that although it is an ultrasonic wave, it carries
optical absorption information. PA waves are generated from within the tissue and
propagate outwards towards the medium surface. At the surface, the PA echoes are detected
through ultrasonic piezoelectric transducers. They are then processed to create an image
which maps the optical absorption distribution of the medium.
Current available technologies available suffer from either poor penetration depth,
as in optical imaging techniques or poor contrast resolution, as in ultrasonic imaging.
Optical imaging such as optical coherence tomography (OCT) can provide excellent spatial
resolution (Huang 1991). However, its penetration depth is severely limited as a result of
the strong diffusivity of light biological tissue (Fercher et al. 1997). Also, while optical
imaging techniques are sensitive to the light backscattering that is related to tissue
2

morphology, they are insensitive to the optical absorption that is related to important
biochemical information. High-frequency ultrasonic imaging (US) can achieve good
resolution as well as deep image depth, but has a poor contrast determined by weak acoustic
backscatter, falling short of its optical counterparts. The varied optical absorption of states
of hemoglobin allows PA techniques to also extract some functional imaging information
(Zhang et al. 2006). PA imaging can essentially bridge the gap between technologies by
providing optically based contrast with an ultrasonic resolution at greater depths. Table 1-1
provides a comparison of US, OCT, and PA imaging.

Table 1-1: Comparison of low-frequency US, high-frequency US, OCT, and PA imaging.
Modality Lateral Resolution (μm) Contrast
Imaging Depth
(mm)
Low-frequency US (< 10 MHz) > 200 (Poor) Fair > 100 (Excellent)
High-frequency US (> 10 –
50MHz )
80-200 (Good) Fair 3 – 15 (Good)
OCT ~15 (Excellent) Excellent 1.5 (Poor)
PA imaging
80-200 (No Speckle,
Good)
Good 3 – 10 (Good)

In this dissertation, novel high-frequency ultrasonic transducer designs have been
investigated for improving the performance of PA imaging, including employing new
advanced materials, progressing to a higher frequency, better sensitivity, and specifically
a proper configuration for the efficient light delivery. It is hoped that these designs will
yield better resolution, deeper penetration depth, and improved contrast for PA imaging.
Since PA imaging is a hybrid modality, a review of both the photoacoustic and
ultrasonic transducer theory is included in this thesis. Transducer design modeling
3

softwares used in this research and introduces several standard test procedures to evaluate
transducer performance. (Chapter 2 and Chapter 3). Chapter 4 describes the development
of high-frequency highly focused ring shape transducer for high resolution photoacoustic
microscopy (PAM) application. Chapter 5 and Chapter 6 present the development of
miniaturized high-frequency single element transducers and phased array transducers for
photoacoustic endoscopy (PAE) application. Chapter 7 summarizes the new achievements
on the two areas and discusses future work as well.

4

Chapter 2 Photoacoustic Theory
2.1 Photoacoustic Effect
The PA effect, also referred as optoacoustic effect, was first observed by Alexander
Graham Bell in 1880 (Bell 1880). He found that absorption of electromagnetic waves by
medium generated sound waves. Even though the PA effect was known for a long time, it
was not applied in highly scattering media until 1994 by Kruger (Kruger and Liu 1994). A
few years later it was applied in biomedical imaging (Hoelen et al. 1998).
The basic principle of the PA effect is schematized in Figure 2-1. Light is shone
on a sample that absorbs a fraction of the incident energy. This energy is converted into
heat. The temperature rise caused thermoelastic expansion of the object. This sudden
pressure rise generates sound wave propagating to the surrounding medium which can be
detected. By detecting the pressure waves using an ultrasonic transducer, where light was
absorbed can be localized and obtain important information about the studied sample.
5

Figure 2-1: Basic principle of the PA effect.
2.2 Photoacoustic Imaging
The PA effect explains how electromagnetic energy can be absorbed and converted
into acoustic waves. PA imaging benefits from the advantages of pure optical or ultrasound
imaging, without the major disadvantages of each technique (Ntziachristos et al. 2005). PA
imaging combines the high contrast from absorption of light with the high resolution and
deep penetration depth of ultrasound imaging.
In PA imaging, laser pulses are delivered into biological tissues. The delivered
energy will be absorbed and converted into heat, inducing transient thermoelastic
expansion and thus wideband ultrasonic emission. The generated ultrasonic waves are
detected by ultrasonic transducers to form images. It is known that optical absorption is
closely associated with physiological properties, such as hemoglobin concentration and
oxygen saturation (Grinvald et al. 1986). As a result, the magnitude of the ultrasonic
6

emission (i.e. PA signal), which is proportional to the local energy deposition, reveals
physiologically specific optical absorption contrast.
Pure optical imaging like OCT can provide excellent images for very superficial
structures. However, it is limited by its short penetration depth of no more than 2 mm
(Welzel et al. 1997). Unlike OCT, PA imaging is diffraction limited, meaning limited by
the frequency of the detected PA signals and not by optical diffusion. This technology does
not depend on single backscattered light as OCT does. In contrast to OCT, the scattering
of light works as an advantage in PA techniques because any light, including both singly
and multiply scattered photons, contributes to the tissue absorption and improves tissue
illumination homogeneity (Niederhauser et al. 2005). The imaging depth of PA imaging is
also limited by the ultrasonic attenuation which will be described in Chapter 3.
Consequently, the imaging depth capability of PA imaging far exceeds that of OCT.
Because PA waves travel one-way to reach the ultrasonic transducer, PA imaging
is essentially speckle free. Just as in ultrasound tomography, temporally detected
transmission echoes are stronger than reflected or backscattered echoes. The PA wave is
representative of the optical absorption information of the medium, but is also comprised
of an acoustic one-way traveling echo. The acoustic portion of the traveling wave is still
subject to the acoustic effects, such as random scattering of wavelength and smaller sized
particles. However, the acoustic wave is generated from within the tissue. This means it is
a one-way traveling wave that carries mainly optical contrast information, not ultrasonic
backscatter. Therefore, speckle resulting from the acoustic properties of the medium does
not offer a significant contribution to degrade the image, and the dominant contrast
mechanism is optical contrast.
7

High-frequency US imaging has good resolution for the image depth achieved, but
is limited by a weak image contrast. PA imaging combines the contrast advantage of optical
imaging with the resolution advantage of US imaging. The contrast of the reconstructed
PA image is related to the optical properties of the tissue, while the resolution is not limited
by optical diffusion or photon scattering like pure optical imaging. Natural sources for
strong optical contrast in biological tissue include hemoglobin and other chromospheres
present in the tissue (Beard and Mills 1997). In addition to structural visualization, the
varying absorption of different states of hemoglobin allow for functional imaging
capabilities of PA imaging.
2.3 PA Imaging Modalities in Biomedicine
Currently, there are three main approaches of PA imaging modalities: 1.
photoacoustic tomography (PAT) (Wang et al. 2003), 2. photoacoustic microscopy (PAM)
(Zhang et al. 2006), and 3. photoacoustic endoscopy (PAE) (Yang et al. 2009). The method
of acoustic generation is the same for all three techniques. A Q-switched pulsed laser is
used as the electromagnetic source to irradiate tissue. This setup begins with a pumping
laser (usually a Nd:YAG type), which provides a continuous wave beam at a specific
wavelength. The Q-switch is essentially an aperture which allows the light to accumulate,
then releases it, sending out a burst. The light can be coupled to a fiber or directed through
a lens at the target medium. The PA waves thereby generated are detected by the ultrasonic
transducer or transducer array, and traverse receiving, signal processing, and digitization
steps before the electromagnetic absorption distribution is reconstructed to create the image.
8

A typical PAT system uses an unfocused ultrasound transducer to acquire the
photoacoustic signals, and the image is reconstructed by universal back projection
algorithm (Xu and Wang 2005). PAM and PAE systems, on the other hand, use a
spherically ultrasound detector with two dimensional point-by-points scanning and require
no reconstruction algorithm.
The spatial resolution of PAT is determined by ultrasonic detection in the PA
emission phase (Wang 2008). The center frequency and the bandwidth of the ultrasonic
detection predominantly determine the spatial resolution of PAT. The greater the center
frequency and the wider the bandwidth, the better the spatial resolution is.
Unlike PAT, which depends on reconstruction algorithms to form images, PAM,
by focusing both the optical excitation and ultrasonic detection, detects PA waves coming
primarily from the focal zone (Oraevsky and Karabutov 2003). The dual focal zones are
usually configured confocally to maximize sensitivity (Wang and Hu 2012). The axial
resolution is determined by the acoustic time of flight, whereas the lateral resolution is
determined by the overlap of the dual focal zones. Depending on whether the optical or
ultrasonic focus is finer, PAM is further classified into optical-resolution PAM (OR-PAM)
(Maslov et al. 2008; Hu et al. 2011) and acoustic-resolution PAM (AR-PAM) (Zhang et al.
2006).

9

Chapter 3 Ultrasonic Transducer
3.1 Ultrasonic Transducer Fundamentals
3.1.1 Piezoelectric Parameters
Transducers are the key devices which make ultrasonic imaging possible. The
image quality is affected by physical and electromechanical characteristics of an ultrasound
transducer. Piezoelectric based ultrasonic transducers convert electrical energy to
mechanical energy (acoustic waves), and vice versa via a phenomenon called piezoelectric
effect. When an electrical signal is applied across the material it forces the electric dipoles
to realign and thus, produces a physical change in thickness. Likewise, when a mechanical
force such as an ultrasonic pressure wave is applied, there is a resultant electrical potential
generated (Shung 1996).
The criteria of a high quality ultrasonic transducer for imaging applications involve
broad bandwidth or wide frequency response, good impedance matching to biological
tissues, high efficiency as a transmitter, and high sensitivity as a receiver. The aperture size
of the transducer is also very critical. Piezoelectric material plays a crucial role in
fabricating the ultrasonic transducers, since it is the only active material inside the
transducers. The properties of the piezoelectric materials are mostly determined by
electromechanical coupling coefficient (kt), acoustic impedance (
a
Z ), dielectric constants,
and clamped dielectric permittivity (
33 0
/
s
 ).
10

Piezoelectric materials with high
t
k value are efficient in energy conversion,
suggesting the improved sensitivity of the transducers. The value of
t
k can be calculated
by equation 3-1:
tan( )
22
ar r
t
aa
ff f
k
ff
 
 3-1
where
r
f is the series resonant frequency at which the conductance reaches the minimum
and
a
f is the parallel resonant frequency at which the resistance reaches the maximum.
The piezoelectric constants relating the mechanical strain are termed the strain
constants, or the "d" coefficients. The piezoelectric d coefficients can be related to the
polarization generated per unit of mechanical stress applied to a piezoelectric material
(piezoelectric direct effect), or the mechanical strain produced by a piezoelectric material
per unit electric field (piezoelectric converse effect). The unit can be expressed by coulomb
per newton or meter per volt from the above definitions:
strain development short circuit charge density
applied electric field applied mechanical stress
d 3-2
33
d denotes the direction of excitation (force or electric field) and response (charge or
displacement) are along the 3 direction (along the polarization axis). Since the d33
coefficient relates closely to the mechanical displacement performance, a high sensitivity
piezoelectric material should have high
33
d .
The piezoelectric constants relating the electric field produced by a mechanical
stress are termed the voltage constants, or the "g" coefficients. The unit can be expressed
as volt-meter per newton.
11

open circuit electric field
applied mechanical stress
g  3-3
Output voltage is obtained by multiplying the calculated electric field by the thickness of
ceramic between electrodes.
33
g indicates that the electric field and the mechanical stress
are both along the polarization axis. A piezoelectric material with good receiving capability
should have high
33
g .
33
g is related to
33
d by the relation:
33
33
0 33
T
d
g

 3-4
where
0 33
T
 is the dielectric constant measured at 1 kHz.
The acoustic impedance (
a
Z ) of a material is defined as the product of its density
(  ) and longitudinal acoustic velocity (
t
v ):
at
Zv   3-5
t
v is determined by the elastic constant (
33
c ):
33
t
c
v

 3-6
33
c is related to the density (  ), the thickness of the piezoelectric element ( t ), and the anti-
resonance frequency (
a
f ):
2
33
(2 )
a
c tf   3-7
Meanwhile,
33 0
/
s
 is a critical issue considering the electrical impedance
matching of the transducers to the 50 Ω imaging electronics, which would affect the
sensitivity in both transmitting and receiving. The electrical impedance of a transducer (
e
Z )
is inversely proportional to the surface area of the piezoelectric element ( A ) and
33 0
/
s

value (Cannata et al. 2003):
12

33 0 33
1
e TT
t
Z
CA 
 3-8
For a miniaturized transducer, materials with high dielectric permittivity are more desirable.
3.1.2 Piezoelectric Materials
Many ceramics, polymers, and crystals exhibiting superior piezoelectric properties
have been widely used for ultrasonic transducer applications, such as lead zirconate titanate
(PZT), polyvinylidene fluoride (PVDF), lithium niobate (LiNbO3), and lead magnesium
niobate-lead titanate (PMN-PT). Comparisons of major properties of the piezoelectric
materials mentioned above are listed in Table 3-1. Due to the high receiving constant
33
g
and low acoustic impedance (close to the medium), PVDF transducers are commonly used
in PA imaging for having high receiving performance and wide band frequency response
(Wang et al. 2003). However, the PVDF is not suitable for being high sensitivity transducer
element because of its low
t
k and
33
d . Moreover, it cannot be used to fabricate transducers
with small aperture size because it has very low
33 0
/
s
 . LiNbO3 single crystal
simultaneously possesses high
t
k and low
33 0
/
s
 . This feature makes it possible to build a
transducer with high sensitivity, large bandwidth, and large aperture size. Among all
piezoelectric materials, PMN-0.33PT single crystal is a promising candidate for
miniaturized highly sensitive transducer design with high kt ,
33 0
/
s
 and d33.

13

Table 3-1: Properties of important piezoelectric materials.
Property
PZT-5H
(Zipparo, 1997)
PVDF
(Bloomfield,
2000)
36
o
Y-cut
LiNbO 3
(Cannata, 2003)
PMN-0.33PT
(Zhang, 2001)
Type
Fine grain
Ceramic
Polymer Single crystal Single crystal
t
k 0.51 0.13 0.49 0.58
12
33
(10 / ) d C N

593 -33 6 1500
3
33
(10 / ) g v m N

 19 -373 23 29
33 0
/
s
 1470 6.5 28 680
3
( / ) g cm  7.5 1.8 4.64 8.06
( / ) v m s 4580 2150 7340 4610
()
a
Z Mrayl 36.0 3.87 34.0 36.9

3.1.3 Mechanical Matching
As discussed previously, in ultrasound imaging applications, a transducer with a
broad bandwidth and low loss is preferred. Mechanical matching, which includes front
matching and rear backing, is a well-established solution to a high performance ultrasonic
transducer.
A simplified transducer model is illustrated in Figure 3-1. When a single electrical
pulse is applied across the piezo-electrical element, pressure waves a and b will be
produced with opposite directions on the front and the rear surfaces of the element,
respectively.
14

Figure 3-1: A simplified model of an ultrasonic transducer.

The pressure wave b moves forward and reaches the front surface of the piezo-
element. The pressure transmits into the front loading medium at a normal incidence is
governed by the transmission coefficient:
2
l
pl
Z
T
ZZ


3-9
where
l
Z and
p
Z are acoustic impedance of the loading medium and piezo-elements,
respectively. The acoustic impedance
p
Z of a typical piezo-element (normally around 30
MRayl) is much higher than the acoustic impedance of the loading medium, such as human
tissues, which is around 2 MRayl. With the typical values, the transmission coefficient T
is calculated to be only around 10%. Therefore, a matching layer was inserted in between
piezoelectric materials and the loading medium to compensate for their acoustic impedance
mismatch. According to transmission line theory, 100 % transmission occurs when
thickness of the matching material is equal to
4
m

(
m
 is the wavelength of the matching
material) and acoustic impedance of the matching material
m
Z satisfies:
15

12
()
m p l
Z Z Z  3-10
However, for broadband transducers application, the above equation should be modified to
(Desilets et al. 1978):
2 1 3
()
m p l
Z Z Z  3-11
Sometimes, when the acoustic impedances of the piezoelectric materials are very
high, two quarter-wavelength matching layers are suggested (Cannata et al. 2003). Desilets
et al showed that the acoustic impedances of the two matching layers should be respectively
equal to (Desilets et al. 1978):
7 / 1 3 4
1
) (
l p m
Z Z Z  3-12

6 1/7
2
()
m p l
Z Z Z  3-13
Popular acoustic matching material includes silver epoxy (7.3 MRayl) and parylene (2.3
MRayl), which are close to the above requirements.
Similarly, when the pressure wave a moves backward and hits the rear surface of
the piezo-element at a normal incidence, a fraction of energy will be transmitted into the
air and the rest is reflected back and is governed by the reflection coefficient:
ap
ap
ZZ
R
ZZ



3-14
where
a
Z is the acoustic impedance of the air, which is close to zero. It is obvious that a
majority of the energy will be reflected back to the front face. In reality, even with matching
layers, 100% transmission is impossible at the front surface. The reflected wave thus will
reverberate inside the piezo-element, causing a long ring (narrow bandwidth) of the
produced ultrasonic pulse. Backing, therefore, is used to damp out the ringing due to
acoustic impedance mismatch between the air and piezoelectric materials. Ideally, when
16

the acoustic impedance of the backing (
b
Z ) is equal to the piezo-element (
p
Z ), there will
be no reflection and a monocycle pulse can be generated. However, the sensitivity will be
significantly reduced in this case. A compromise has to be made between bandwidth and
sensitivity in practical applications. The transducers fabricated in this research are mainly
using conductive epoxy, E-solder 3022 (Von Roll Isola Inc., New Haven, CT). This
material has a high attenuation (120 dB/mm at 30 MHz) and relatively low acoustic
impedance (5.9 MRayl), making it possible to achieve low insertion loss, short and well-
shaped pulses suitable for imaging purpose. Besides lossy, the backing materials are
usually rigid so as to support the fragile piezo-elements.
3.2 Single-element and Array Transducers
3.2.1 Single-element transducer
Single-element transducer is the most elementary ultrasonic transducer. A
broadband single-element transducer normally consists of a piezo-element which was
sputtered with electrodes on both surfaces, matching layers, backing material and
sometimes a lens. The functions of the backing and matching layers have been explained
in the previous section. The lens here is used to focus the ultrasonic beam to a desired
distance. For a piston transducer with radius of a and wavelength of λ in the loading
medium, there is a distance which is the last maximum of the axial pressure. The distance
was found to be:
2
0
a
Z

 3-15
17

The region between transducer surface and this distance is called near-field zone or
Fresnel zone. In this zone, the axial pressure oscillates. The region beyond the distance is
called far-field zone, where the axial pressure decreases gradually. The far-field region is
the ultrasound imaging region. This means the target of interest should be always located
beyond
0
Z .
In the far-field, directivity function of a piston transducer is found to be:
1
2 ( sin )
()
sin
J ka
H
ka



 3-16
where
1
J is the first order of the Bessel function. With the equation, the main lobe shape
can be related to:
sin 0.61
a

  3-17
For a rectangular element which is the basic unit of an array, its directivity function
is:
sin sin
sin sin
2 2
( , )
sin sin
2 2
x
xy
x
kc kb y
H
kc kb y
 



   
   
 3-18
And its main lobe angel can be related to:
11
sin , sin
xy
cb



 3-19
where c and b is the dimension c in x-direction and dimension b in y-direction
respectively.
18

3.2.2 Array Transducer
To produce a two-dimensional US image by a single-element transducer,
ultrasound beam needs to be scanned across the imaging plane by mechanical scanning.
Unlike single-element transducers, arrays have the advantage of obtaining multiple scan
lines by electronically scanning without mechanical repositioning to create an image.
Transducer arrays are the current and future trend for ultrasonic imaging. Arrays employ
many piezoelectric elements, where each element has its own electrical connection.
Elements can be excited individually or in groups, effectively forming a sub-aperture. They
provide two enormous advantages over single element transducers: the ability to
electronically steer and electronically focus a beam. Transducer arrays create a system
where different delays can be assigned to the received echoes from each element by
beamforming, and allowing for an adjustable electronic focus without mechanical scanning.
Linear and phased arrays are two main commonly used transducer array types. Linear
arrays form scan lines by using a sub-aperture of elements that slides across the face of the
array. Each sub-aperture forms a single scan line, and is then translated by one array
element to the next position, producing a rectangular image field. On the other side, phased
arrays generate angular scan lines by using all elements at once to steer the main beam at
different angles.
Like a single-element transducer, a backing material and one or two matching
layers are used to improve arrays’ performance. To minimize acoustic cross-talk,
piezoelectric material is diced into small elements. The space between two elements (g) is
called a kerf. The kerfs may be filled with acoustic isolating material. The isolating material
serves can decrease acoustic cross-talk and provides a rigid support for array elements.
19

Sometimes, a lens is used. It can focus the acoustic beam in the elevation direction thus to
decrease slice thickness of the imaging plane which can cause serious image artifacts; at
the same time, it can serve as a protecting layer for the fragile array elements.
The radiation pattern of a linear array in the far field is given by:
 
1
sin sin
N
m
bu Lu
H u c u m c
g





   
   
    
   


3-20
where
sin
x
u  
,   Hu is the direction function at an angle of
x

, L is length of the array,
N is the number of the elements, b is the width of the element, g is pitch which is the
space between the center of two adjacent elements,  is the wavelength in the loading
medium.
The linear array radiation pattern equation indicates that at certain angles big side
lobes called grating lobes may occur. The angle where the grating lobe appears is governed
by the following equation:
1
sin
g
n
g






3-21
where n is an integer. Therefore, the pitch should be less than  to make sure the grating
lobes occur at angles larger than 90° (in a practical design, g is normally equal to 0.75 –
2  ). With the above two equations, other rules for designing a linear array can also
be obtained: the width of the element should be as large as possible in order to damping
the magnitude of the grating lobes; the aperture of the array should be as large as possible
in order to get a narrower main beam width; and ratio of the width to thickness of
the element (
b
t
) should be < 0.6 to avoid lateral resonance. On the other side, for phased
arrays, the pitch g should be smaller than 0.5  ,
b
t
< 0.6, and
2
b

 to ensure a
20

broad beam because the beam is steered. However, above requirements can’t be satisfied
at the same time, some trade-offs have to be made in a practical design.
Phased arrays allow dynamic focusing and beam steering. Dynamic focusing can
be achieved in transmission and in reception. However, multiple transmissions of pulses
are needed for dynamic focusing during transmission, slowing down the frame rate.
Transmission dynamic focusing is usually done in discrete zones, whereas receiving
dynamic focusing can be done in many more zones or almost continuously.
3.3 Axial and Lateral Resolution
Image resolution is important in any imaging modality. Axial resolution refers to
the minimum reflector spacing along the axis of a beam that results in separate,
distinguishable echoes. It is the spatial extent of the echo in the axial direction and is mostly
determined by pulse duration. Lateral resolution refers to the ability to distinguish two
closely spaced reflectors positioned perpendicular to the beam axis (Lockwood et al. 1996).
It is more concisely defined as the spatial extent of the beam profile in the lateral direction
within -3dB or -6 dB of the peak intensity. Basic imaging theory has established that lateral
resolution is a function of the wavelength of sound, the depth of the scan, and the size of
the aperture used to form an image. A common measurement of the aperture size is f
number (
#
f
), which is defined as the ratio of focal depth to aperture size.
Ultrasound refers to the sound with frequency above the human hearing range
(20~20,000 Hz). High-frequency ultrasound refers to ultrasound with frequency above 15
21

MHz. The working frequency directly determines the spatial resolutions of an image,
described in equation 3-22 and 3-23.
#
0
lateral
c
Rf
f

3-22

6
2
axial
dB
c
R
BW




3-23

where C is the speed of sound, f0 is the center frequency of transducer, f# is defined as the
ratio of focal distance to the aperture diameter of transducer, and BW-6dB represents the -6
dB bandwidth of transducer. Equation 3-23 is valid only for a focused transducer in the
focal zone. It is obviously that increasing the center frequency will improve both axial and
lateral resolutions.
3.4 High-frequency Ultrasound Imaging
Ultrasound is a compressional wave that can only propagate inside a medium. Once
the transmitting ultrasonic waves hit an interface, where acoustic impedances differ in the
mediums on the two sides, parts of the ultrasonic waves will be bounced back and received
by the same transducer. The rest of ultrasonic waves continue propagate until hit other
interfaces. The position of each interface can be timely resolved according to equation 3-24,
where d is the distance between transducer surface and interface of acoustic impedance
discontinuity, c is sound speed in the medium, and t is the time delay of echo signal.
2
tc
d  3-24
The echo signals received along one transmitting/receiving route is named one A-
line. If the ultrasonic transducer is physically scanned linearly or rotationally, a 2-D image
22

can be formed by incorporating multiple A-lines. Echo signal strength is mapped to the
gray scale in the image.
3.5 Issues of Ultrasonic Attenuation at High-frequency
As mentioned in Chapter 1 and Chapter 2, the imaging depth of PA imaging is
limited by the ultrasonic attenuation. Although high-frequency ultrasound is able to
provide excellent spatial resolutions, a major concern is the high attenuation effect. The
attenuation of acoustic energy is caused by reflection, scattering at interfaces, and
absorption. The pressure of a plane wave () Pz traveling in the Z direction displays the
following exponential behavior (Shung 1992) in equation 3-25.
0
()
z
P z P e
 
 3-25
Where P0 is the ultrasonic pressure at transducer surface, α is the acoustic attenuation
coefficient of the medium and z is the traveling distance. Attenuation coefficient α is
frequency dependent and also has strong dependence on the type of tissues. It can be
defined as


  f
0
 3-26
where
0

is the attenuation coefficient at 1 MHz and γ is the frequency dependence
parameter. Attenuation coefficient from 10 to 100 MHz was summarized and plot in Figure
3-2.
23

Figure 3-2: Attenuation of variety tissues in the frequency range from 10 to 100 MHz
(Foster et al. 2000)
3.6 Ultrasonic Transducer Modeling and Testing
3.6.1 Krimholtz, Leedom, and Matthaei (KLM) equivalent circuit model
Popular 1-D transducer models include Mason model (Mason 1948; Mason 1958),
the Redwood model (Redwood 1961) and KLM model (Krimholtz 1970). Among them,
the KLM is more physically intuitive, thus is widely used. The thicknesses and dimension
of each matching layer and piezoelectric layer could be simulated based on this model,
which helps optimize each parameters to achieve the optimal performance of a transducer.
This model divides a piezoelectric element into two halves, each represented by an acoustic
transmission line. The acoustic transmission line serves as a secondary circuit, which is
linked with an electric primary circuit by an ideal transformer as shown below in Figure
3-3 (Shung 2006).
24

Figure 3-3: The KLM model for a single-element transducer.

The component and constant used in the model are list below:
A = transducer area
d = thickness of the piezoelectric material
 = density of the piezoelectric material
c = the speed of sound in the piezoelectric material
c
Z
= radiation impedance of the piezoelectric layer (=ρcA)
*
 = complex clamped dielectric permittivity
33
e
= piezoelectric constant
D
c = elastic constant of piezoelectric layer
s
r
 = clamped dielectric constant
*
0
A
C
d


3-27
2
2
33
r
Ds t
e
k
c 
 3-28
25

0
2 c
d

 
3-29
00
t
c
k
CZ


 3-30
0
21
0
(sin ( ))
t
C
C
kc 



3-31
With the KLM model, the electric impedance of the transducer is given as:
12
2
0
12
11
()
in
ZZ
Z
j C j C
ZZ

  

3-32
where
1
Z
and
2
Z
are the input impedances of the acoustic transmission line looking
towards the front acoustic port and back acoustic port, respectively. Equations that describe
insertion-loss and pulse-echo response et al can also be obtained with the KLM model.
From the design point of view, the KLM model allows an intuitive approach to be used in
optimizing the transducer performance: the two acoustic ports can be used to interpret front
matching and rear backing; the electrical port, on the other hand, can be used to explain
electrical matching.
3.6.2 Transducer Modeling Softwares
During transducer design process, three transducer modeling softwares were used
to help optimize transducer’s design including PiezoCAD (Sonic Concepts, Inc. Bothell,
WA), Field II (Jensen 1992; Jensen 1996), and PZFlex (Weidlinger Associates, Inc.,
Mountain View, CA).
PiezoCAD is a piezoelectric transducer modeling software package based on KLM
model to calculate the overall transducer characteristics from the electric terminals to the
26

front acoustic port. The user can select piezoelectric material parameters from extensive
piezoelectric database tables, including plate, beam, and bar mode elements in ceramic,
crystalline, polymer, and composite materials. The user can enter multiple acoustic
matching and backing layers from acoustic database tables, as well as load medium
characteristics. Thickness, velocity, attenuation, and cross-sectional area can be entered for
each layer. On the electrical port, the user can specify a matching network including any
combination of series or shunt resistors, inductors, or capacitors; transformers; coaxial
cable, and multiple identical piezoelectric layers that are electrically connected in parallel.
PiezoCAD output includes a selection of acoustic and electric input immittance functions;
transmit and receive power conversion efficiencies; transmit, receive, and pulse-echo
transfer functions and time domain waveforms.
Field II is a program for simulating ultrasound tranducer fields and ultrasound
imaging using linear acoustics. The program uses the Tupholme-Stepanishen method for
calculating pulsed ultrasound fields. The program is capable of calculating the emitted and
pulse-echo fields for both the pulsed and continuous wave case for a large number of
different transducers. Also any kind of linear imaging can be simulated as well as realistic
images of human tissue. The program is running under Matlab (MathWorks, Natick, MA).
Until recently, medical transducer designers relied almost exclusively on 1D
analytical models and experimental prototypes. Now, many employ comprehensive finite
element simulations for transient, 2D and 3D analyses. Finite element modeling is being
adopted in the design of ultrasonic transducers and imaging arrays. Based on finite element
analysis, PZFlex is widely used and is a useful tool for designing transducer.
27

3.7 Transducer Performance Evaluation
The ultimate indication of ultrasonic transducer performance is its ability to form
an image. However, image quality depends not only on the transducer but also on system
electronics and signal processing algorithms. Hence it is difficult to compare different
transducers only based on images they generated. Therefore, several standard non-imaging
transducer tests were performed for the transducers built in this research.
3.7.1 Electrical Impedance Test
The electrical impedance magnitude and phase angle for each element were
measured using an impedance analyzer (HP 4294A, Agilent Technologies, Inc., Santa
Clara, CA) with the impedance probe (HP 42941A, Agilent Technologies, Inc., Santa Clara,
CA). For this test the transducer face was placed in a water bath, and the electrical
impedance magnitude and phase angle were measured over the transducer pass-band.
3.7.2 Pulse-echo Test
The pulse-echo response test if the most common test performed on ultrasonic
single-element and array transducers. This test was used to measure and provide transducer
center frequency, bandwidth, focal depth, pulse length, and sensitivity. For measuring the
pulse-echo response, the transducer was mounted on a holder and immersed in a tank filled
with the distilled water. The flat quartz reflector was placed at certain distance away from
28

the transducer surface, which is the focal length of the transducer. The set-up of pulse-echo
test is shown in Figure 3-4.

Figure 3-4: Set-up of pulse-echo test.

By connecting to an ultrasonic pulser-receiver (Panametrics 5900PR, Olympus
NDT Inc., Waltham, MA), the transducer were excited by a 1 μJ electrical impulse with
200 Hz repetition rate and 50 Ω damping factor. The echo signals were acquired and
displayed using LC534 1GHz digital oscilloscope (LeCroy Corporation, Chestnut Ridge,
NY). The captured pulse-echo response signals were then used to compute the frequency
spectrum by Matlab (R2012b, MathWorks, Natick, MA). The center frequency (fc) and -6
dB bandwidth (BW) of the array were determined from the measured FFT spectrum:
12
2
c
ff
f

 3-33

21
100%
c
ff
BW
f



3-34
where
1
f
and
2
f
represent the lower and upper -6 dB frequencies, respectively.
29

3.7.3 Insertion Loss Test
The two-way insertion loss (IL) or the relative pulse-echo sensitivity is the ratio of
the transducer output voltage
o
V
to the excitation voltage
i
V
delivered to the array from a
driving source. The set-up of insertion loss test is shown in Figure 3-5. The following
equation 3-35 was used for calculation:
42
20log 1.9 2.2 10 2
o
c
i
V
IL d f
V

      3-35
Where d was the distance (mm) between target and transducer surface. The attenuation
compensation for quartz is 1.9 dB and the attenuation coefficient compensation for water
is
42
2.2 10 / / dB cm MHz

 .
The set-up of insertion loss test is shown in Figure 3-5. The transducer was
connected to a function generator (AFG2020 function generator, Tektronix, Inc.,
Beaverton, OR) which was used to generate a tone burst of 30-cycle sine wave at
c
f
. The
echo signal received by the transducer,
o
V
, was measured by the oscilloscope with 1 M 
coupling. The amplitude of the driving signal
i
V
was then measured with 50  coupling.

Figure 3-5: Set-up of insertion loss test.

30

3.7.4 Crosstalk Measurement
For array transducer, the level of electrical and acoustical separation between
elements was determined by measuring crosstalk. The set-up of crosstalk measurement is
shown in Figure 3-6. For this test the array was positioned in a degassed/deionized water
bath opposite an absorptive piece of rubber. A Sony/Tektronix AFG2020 function
generator (Tektronix, Inc., Beaverton, OR), set in burst mode, was used to excite a
representative element with the applied voltage measured as a reference. Voltages across
the three nearest elements were measured and compared to this reference voltage. This
process was repeated at discrete frequencies over the array pass-band.

Figure 3-6: Set-up of array crosstalk measurement.

3.7.5 Directivity Measurement
The set-up of array directivity measurement is shown in Figure 3-7. A needle
hydrophone (Precision Acoustics, Dorchester, UK) was placed at the elevation focus and
31

used to acquire the amplitude of the time-domain response from a array element. For array
transducer, the azimuthal one-way directivity was measured by rotating a representative
element around an axis along its length and center and comparing received signal amplitude
received by hydrophone at discrete angular positions.

Figure 3-7: Set-up of array directivity measurement.

32

Chapter 4 Highly Focused High-frequency Ring Transducer
for Photoacoustic Microscopy Application
4.1 Introduction to PAM
In order to determine cancer malignancy, pathologists first must find the malignant
cells in cancer lesions (Connolly 2010). Imaging of cell nuclei is very critical to cancer
diagnosis. The nuclei in cancer cells have typical morphological features, such as irregular
shapes and large sizes (Zink et al. 2004), which allow pathologists to identify cancer cells
through microscopic examination. Optical microscopy of cell nuclei is the primary
histological method, widely used for cancer diagnosis and malignancy grading (Lester
2010). Ultraviolet photoacoustic microscopy (UV-PAM) is a new technique capable of
specifically imaging of cell nuclei (Yao et al. 2010). UV-PAM produces in vivo images of
unstained cell nuclei with specific, positive, and high-image contrast.
Figure 4-1 shows the set-up of a conventional PAM system. Conventional PA
systems have complex and thick hand made parts in a path of the optical objective and
complicated acoustic path which causes losses and reverberations. Optical absorption in
prisms, glue, lens and medium is also a main issue to be considered.
33

Figure 4-1: Set-up of a conventional PAM system.

4.2 PAM Transducer
To overcome the drawbacks of conventional PAM system, a special PAM (UV-
PAM) system was developed. The details of the UV-PAM system will be described in
Chapter 4.3. For the new system, a ring-shape transducer was designed, as shown in Figure
4-2. By placing optical components at the hole, the ring transducer provides coaxial
alignment of optical light pulses delivery and ultrasound detection. This configuration can
enhance the performance of PAM system in many aspects such as capability of using any
optical wavelength, minimization of size of the system, convenient system implementation,
light delivery simplification, better axial contrast, and aberration free image.

34

Figure 4-2: Schematic of PAM transducer.
4.2.1 Transducer Design
The inner diameter of the transducer is designed at 2 mm which is wide enough for
the light go through and focus at the desired focal distance. To achieve large aperture size
desired for the PAM transducer, lead-free single crystal lithium niobate (LiNbO3) was
selected as piezoelectric material due to its low dielectric permittivity. To improve the
sensitivity and bandwidth of the transducer, two matching layer strategy (Cannata et al.
2003) was employed for the transducer. The transducer is designed with a small
#
f
to
achieve high lateral resolution. Before fabrication, PiezoCAD was used to determine the
proper thicknesses of the lithium niobate, the 1st matching layer and the 2nd matching
layer.

Table 4-1: Design parameters and modeling results of PAM transducer.
Specifications Values
Center frequency 50 MHz
Bandwidth > 50%
Piezoelectric material (PM) LiNbO 3
Thickness of PM 60 µ m
1
st
matching layer (ML) Silver epoxy
Thickness of 1
st
ML 8 µ m
2
nd
ML Parylene
Thickness of 2
nd
ML 9 µ m
35

Backing E-Solder 3022
Thickness of Backing 2.0 mm
Inner Diameter (ID) 2.4 mm
Outer Diameter (OD) 4.8 mm
Focal Distance 5.6 mm
f# 1.17
Center frequency 50.8 MHz
-6dB Bandwidth 60.2%
-6dB Pulse width 31 nsec
Peak Amplitude -47.48 dB re 1 V/V

The parameters of the 50 MHz lithium niobate ring transducer used for simulation
are shown in Table 4-1. Electrical impedance and phase modeling results (Figure 4-3)
showed that the electrical impedance at a phase peak is 47.4 Ω at 58.6 MHz which is very
close to 50 Ω, the series and parallel resonant frequencies are 52.9 MHz and 65.9 MHz,
respectively. Pulse-echo impulse response modeling results of a ring transducer are shown
in Figure 4-4. The simulated -6 dB bandwidth of the transducer is about 50 %.

Figure 4-3: Electrical impedance and phase modeling results of PAM transducer by
PiezoCAD.
36

Figure 4-4: Pulse-echo impulse reponse and its FFT spectrum modeling result of PAM
transducer by PiezoCAD.

To conform the performance of the ring geometry, the influence of the transducer
with an inner hole has been investigated and modeled using Field II. Figure 4-5 shows (a)
an aperture of the PAM transducer illustrated in Field II and (b) a simulated acoustic field
transmission with a contour representation of the transducer. The simulation result shows
that the focused ring transducer has very strong acoustic field near the focus area.

37

(a) (b)
Figure 4-5: (a) The aperture of PAM transducer; (b) The simulation of the acoustic
field transmission of PAM transducer with a contour representation.

4.2.2 Transducer Fabrication Process
A 36º rotated Y-cut. LiNbO3 (Boston Piezo-Optics, Medway, MA) wafer was first
lapped down to 60 µ m and sputtered with Cr/Au (500Å/1000Å) layers as electrodes at top
and bottom surfaces. A silver epoxy matching layer made from Insulcast 501 and Insulcure
9 (American Safety Technologies, Roseland, NJ) and 0.5-1 µ m silver particles (Sigma-
Aldrich Inc., St. Louis, MO) was then cured over the top of the wafer and lapped to the
designated thickness of 8 µ m. A conductive backing material, E-solder 3022 (VonRoll
Isola, New Haven, CT), was then applied to the bottom of the wafer and lapped to 2.0 mm.
The acoustic stack was turned down and drilled on a lathe to the proper inner and outer
diameters, and then a 50 Ω coaxial cable was attached to the backing. The machined
acoustic stack was housed in a thin brass tube. Insulated epoxy (EPO-TEK 301, Epoxy
Technology, Inc., Billerica, MA) was filled into the gap between the acoustic stack and the
-4
-2
0
2
4
-4
-2
0
2
4
0
5

x [mm]
y [mm]

z [mm]
-40
-40
-40
-40
-40
-40
-40
-40 -40
-40
-33
-33
-33
-33
-33
-33
-33
-33
-33
-33
-33
-33
-33
-33
-33
-27
-27
-27
-27
-27
-27
-27
-27
-27
-27
-27
-27
-21
-21
-21
-21
-21
-21
-21
-15
-15
-15
-15
-12 -9
-6
-3
Distance (mm) Along Beam Axis
Distance (mm) Perpendicular to Beam Axis

-3 -2 -1 0 1 2 3
0
1
2
3
4
5
6
7
8
9
10
11
38

brass housing to insulate the inner electrode. The transducer was then focused by hot
pressing a ball bearing of radius equal to the desired focal distance. The ground lead of the
cable was connected to the brass ring by carefully painting on a layer of E-solder 3022.
Another Cr/Au electrode was sputtered over the silver epoxy matching layer and brass
housing to form the common ground connection. A 9-µm-thick parylene layer (± 0.5µ m)
was vapor-deposited onto the transducer and brass housing to serve as second matching
and protecting layer. The transducer’s cable with a SMA type connector was finally
connected to the transducer for electrical connection. The finished transducer is shown in
Figure 4-6.

Figure 4-6: Photo of a PAM transducer.

4.2.3 Transducer Performance Characterization
Measured electrical impedance and phase of the PAM transducer are shown in
Figure 4-7. It was found that the electrical impedance at a phase peak is 50.8 Ω at 62.4
MHz. The series and parallel resonant frequencies are 56.0 MHz and 67.6 MHz,
39

respectively. The values are comparable to the modeling prediction. Measured pulse-echo
waveform and frequency spectrum of the transducer are shown in Figure 4-8. The center
frequency, -6 dB bandwidth, and insertion loss of the transducer were found to be 50 MHz,
88%, and 21.7 dB, respectively.

Figure 4-7: Measured electrical impedance and phase of PAM transducer.

Figure 4-8: Pulse-echo measurement of PAM transducer.

10 20 30 40 50 60 70 80 90
40
65
90
115
140
Frequency (MHz)
Electrical Impedance ( )

10 20 30 40 50 60 70 80 90
-100
-60
-20
20
60
Phase (deg)
Electrical Impedance
Phase
7.3395 7.4648 7.59 7.7153 7.8405
-400
-200
0
200
400
Time ( s)
Amplitude (mV)

10 30 50 70 90
-24
-18
-12
-6
0
Frequency (MHz)
Magnitude (dB)
Pulse-echo Response
Spectrum
40

4.3 UV-PAM System and Imaging
Based on the high-frequency ultrasonic ring-shape transducer, a UV-PAM was
developed for imaging the cell nuclei in intact biological tissue. In UV-PAM, ultraviolet
light (UV) is used as light source instead of visible light. DNA and RNA, two major
compounds in cell nuclei, strongly absorb the UV light around a wavelength of 260 nm
(Stoscheck 1990). In contrast, the UV absorption of protein and lipids is weaker than that
of DNA and RNA by one order of magnitude around 260 nm which is shown in Figure 4-9
(Stoscheck 1990; McHowat et al. 1996; Gill and Vonhippel 1989; Kunitz 1950; Olson and
Anfinsen 1952). To take advantage of this high intrinsic absorption contrast of DNA and
RNA, our UV-PAM system (Figure 4-10) employed light at 266 nm emitted by a Nd:YLF
Q-switched UV laser (QL266-010-O, Crystalaser; pulse width, 7 ns). The laser beam was
spatially filtered by a 25-μm-diameter pinhole (910PH-25, Newport) and focused into a
water tank by a water-immersion objective lens (LB4280, Thorlabs). Subsequently, the
beam passed through the ultrasonic transducer and penetrates a 25-μm thick polyethylene
membrane before the beam focused on the object to be imaged. The polyethylene
membrane was used to seal the bottom of the water tank to form an imaging window while
maintaining acoustic coupling. The laser pulse energy behind the membrane was measured
to be 35-nJ. The specimen was mounted on a two-dimensional scanning stage with a
minimal scan step size of 0.31-μm. Time-resolved photoacoustic signals were detected by
the ultrasonic transducer during raster scanning to reconstruct tomographic images, which
can be rendered in various forms such as cross sectional images and maximum amplitude
projection (MAP) images.
41

Figure 4-9: Absorption spectra of DNA, RNA, and protein.

Figure 4-10: Schematic of the PAM system with a ring-shape transducer.

42

Tomographic images were formed from the amplitude envelopes of the time-
resolved PA signals. Each laser pulse produced a time-resolved PA signal. Hilbert
transformation of the signal produced its amplitude envelope along the z-axis. A collection
of the envelopes along the x-axis produced a cross-sectional image in the x–z plane, a B-
scan image (Zhang et al. 2007). Further scanning along the y-axis produced three-
dimensional images. Projection of the maximal amplitude of each envelope to the scanning
plane (x–y plane) produced a maximum amplitude projection (MAP) image (Zhang et al.
2006; Yang et al. 2009). By using 266 nm light, the UV-PAM system was found to achieve
0.7-μm lateral resolution and 28-μm axial (z-axis) resolution (Figure 4-11).

Figure 4-11: Resolution of UV-PAM.

It took 2.6 minutes to acquire an image of 200 × 200 μm
2
. By scanning with a 1.25
μm step size for 2.6 minutes, the nuclei of the epithelial cells in the mouse lip was imaged.
As shown in Figure 4-12, the image shows a relatively homogeneous distribution of the
cell nuclei, each approximately 6 μm in diameter. The distance between the centers of
neighboring cell nuclei ranges from 16 to 39 μm, suggesting that the stratified squamous
43

epithelium on the lip was composed of cells with a lateral size of the same range. Figure
4-13 shows that the photoacoustic image of cell nuclei matches the histological micrograph.

Figure 4-12: PAM image of epithelial cells in the ex vivo lip of a mouse.

Figure 4-13: (a) Photoacoustic image and (b) histological micrograph of cell nuclei.

Figure 4-14 were acquired at wavelengths of 240, 245, 248, 250, 251, 252, 255,
260, 266, 270, 275, and 280 nm, respectively. Cell nuclei in the mouse skin are shown in
the images at 245, 248, 250, 251, 252, 255, 260, 266, 270, and 275 nm but are unidentifiable
in the images at 240 and 280 nm.

44

Figure 4-14: In vivo en face photoacoustic images of the skin of mouse ears in the form
of maximum amplitude projection (MAP).

Figure 4-15 shows in vivo en face distribution of cell nuclei in the skin of mouse
ears. Figure 4-15(a) is in vivo MAP photoacoustic image of cell nuclei distributed in 1 mm
2

mouse skin, acquired at a wavelength of 250 nm. Figure 4-15(b) is histogram of the nuclear
diameter (n=404). Bin width is 0.6 μm. The solid curve is a Gaussian fit with a mean of
8.6 μm and a SD of 1.6 μm (coefficient of determination R
2
=0.96). Figure 4-15(c) is
45

histogram of the internuclear distance (n=245). Bin width is 2 μm. The solid curve is a
Gaussian fit with a mean of 22.7 μm and a SD of 3.6 μm (R
2
=0.98).

Figure 4-15: In vivo en face distribution of cell nuclei in the skin of mouse ears.

Figure 4-16 shows signal and contrast of in vivo photoacoustic images of cell nuclei
versus the optical wavelength. Figure 4-16(a) is plot of SNR (mean ± SD) of nuclear images
versus wavelength. The SNR was collected from 25 cell nuclei in MAP photoacoustic
46

images for each wavelength. Figure 4-16(b) is plot of CNR (mean ± SD) of nuclear images
versus wavelength. The mean was also calculated from 25 nuclei. The contrast-to-noise
ratio at 250 nm is 2.6 times larger than that at 266 nm.

Figure 4-16: Signal and contrast of in vivo photoacoustic images of cell nuclei versus the
optical wavelength.

Figure 4-17 shows in vivo photoacoustic images of cell nuclei in the skin of mouse
ears with varied laser pulse energy at a wavelength of 250 nm. Figure 4-17(a) shows in
vivo MAP photoacoustic images of cell nuclei at pulse energies of 10, 3, and 2 nJ. Figure
4-17(b) shows CNR (mean ± SD) of the nuclear images. The CNR was collected from 25
cell nuclei in MAP photoacoustic images for each pulse energy.
47

Figure 4-17: In vivo photoacoustic imaging of cell nuclei in the skin of mouse ears with
varied laser pulse energy at a wavelength of 250 nm.
4.4 Summary
A highly focus LiNbO3 ring-shape ultrasonic transducer with 2.4 mm inner
diameter and 4.8 mm outer diameter has been designed and fabricated for UV-PAM
application. The center frequency, -6 dB bandwidth, and insertion loss of the transducer
were measured at 50 MHz, 88%, and 21.7 dB, respectively. The ex vivo PAM image of
micron-level cell nuclei in the lip of a mouse acquired from this transducer demonstrated
the significant improved lateral resolution which achieved submicron level. The UV-PAM
produced images of cell nuclei with specific, positive, and high contrast. The optimal
wavelength is 250 nm for in vivo photoacoustic imaging of cell nuclei. Furthermore, any
wavelength between 245 and 282 nm can be used to produce in vivo images of cell nuclei.

48

Chapter 5 High-frequency Ultrasonic Transducer for
Photoacoustic Endoscopy Application
5.1 Introduction to PAE
Photoacoustic endoscopy (PAE) is a novel technology that embodies photoacoustic
tomography in a small probe that can be used for minimally invasive imaging (Yang et al.
2009; Wang 2008). PAE can provide microscopic or macroscopic depth-resolved cross-
sectional and volumetric images of target tissues (Yang et al. 2009; Oraevsky and
Karabutov 2003). With its unique optical-absorption-based image contrast, PAE compares
favorably to other existing endoscopic imaging modalities, such as Endoscopic ultrasound
(EUS), endoscopic optical coherence tomography (OCT) or confocal endoscopy. EUS is a
medical procedure that uses an endoscope with an ultrasound probe attached to create
detailed pictures of the digestive tract as well as surrounding tissues and organs. Currently,
high-frequency (> 10-MHz) ultrasound has extensively been used in EUS applications
(Foster et al. 2000; Vazquez-Sequeiros and Wiersema 2002; Hurlstone et al. 2005).
However, it suffers from speckle artifacts and poor image contrast. Although OCT and
confocal imaging techniques provide a cellular level (~µ m) image resolution for diseased
tissues in vivo, these techniques are limited by their inability to penetrate deep tissues
because they rely on unscattered, ballistic photons. Conversely, PAE can produce images
with high acoustic resolution and optical-absorption based contrast at depths far beyond
the quasi-ballistic photon regime (Yang et al. 2009; Oraevsky and Karabutov 2003, and
Wang 2009). The high optical absorption-based contrast and depth-resolved image
49

production is attributed to its signal excitation mechanism which utilizes laser pulses
instead of mechanical pulses to generate photoacoustic waves.
At present, single-element side-looking (SL) EUS transducer are commercially
available for use in various diameters (2.0-2.9 mm) and frequencies (12-30 MHz). These
transducers provide high-resolution cross-sectional images. However, since mechanical
scanning is required for the single-element transducers, motion artifacts may be induced to
affect image quality.
The aim of this work is to develop a high-frequency transducer especially for PAE
application without mechanical scanning of probe. To reduce probe size and improve
imaging resolution, three generation of transducers have been developed for this work. The
design, fabrication, and characterization of each transducers will be presented in detail in
this chapter.
5.2 First Generation Transducer Design: Lens-focused Ring Transducer
5.2.1 Transducer Design and Fabrication
To ease the system implementation and minimize the size of the PAE probe, coaxial
alignment of light pulses delivery and ultrasound detection is needed. Consequently, ring
shape ultrasound transducers were designed and fabricated for PAE application.
The US transducer determines the spatial resolution for both PA and US imaging.
Considering the outer diameter of the endoscope and housing thickness, the aperture size
for the US transducer was limited to ~3 mm. To achieve adequate signal sensitivity and
50

high spatial resolution, a lens-focused US transducer was fabricated using the lithium
niobate (LiNbO3) single crystal, and targeted at center frequency of around 40 MHz. Figure
5-1 shows a schematic diagram of the transducer.

Figure 5-1: Schematic diagram of 1
st
generation PAE transducer.

To improve the sensitivity and bandwidth of the transducer, two matching layer
strategy (Cannata et al. 2003) was employed in transducer design. Before fabrication,
PiezoCAD was used to determine the proper dimensions of the piezoelectric material, the
1st matching layer and the 2nd matching layer. The parameters of 40 MHz lithium niobate
ring transducer for simulation are shown in Table 5-1. Electrical impedance and pulse-echo
impulse response modeling results of a PAE transducer are shown in Figure 5-2 and Figure
5-3, respectively. As shown in Figure 5-2, the electrical impedance at a phase peak is 204.7
Ω at 46 MHz, the series and parallel resonant frequencies are 41.2 MHz and 51.4 MHz,
respectively. The simulated -6 dB bandwidth of the transducer is about 55 %.

51

Table 5-1: Design parameters and modeling results of 1
st
generation PAE transducer.
Specifications Values
Center frequency 40 MHz
Bandwidth > 50%
Piezoelectric material (PM) LiNbO3
Thickness of PM 75 µ m
1
st
matching layer (ML) Silver epoxy
Thickness of 1
st
ML 10 µ m
2
nd
ML Parylene
Thickness of 2
nd
ML 12 µ m
Backing E-Solder 3022
Thickness of Backing 2.0 mm
Inner Diameter (ID) 0.7 mm
Outer Diameter (OD) 2.6 mm
Focal Distance 6 mm
f# 2.3
Center frequency 41.1 MHz
-6dB Bandwidth 55.1%
-6dB Pulse width 38 nsec
Peak Amplitude -50.51 dB re 1 V/V

Figure 5-2: Electrical impedance and phase modeling results of 1
st
generation PAE
transducer by PiezoCAD.
52

Figure 5-3: Pulse-echo impulse reponse and its FFT spectrum modeling result of 1
st

generation PAE transducer by PiezoCAD.

Similar to the work described in Chapter 4, Field II was performed to confirm the
performance of the ring geometry. Figure 5-4 shows (a) an aperture of the PAM transducer
illustrated in Field II and (b) a simulated acoustic field transmission with a contour
representation of the transducer. The simulation results show that focused ring transducer
has very strong acoustic field near the focus area.
53

(a) (b)
Figure 5-4: (a) The aperture of 1
st
generation PAE transducer; (b) The simulation of the
acoustic field transmission with a contour representation of 1
st
generation PAE
transducer.

Since the PAE transducer material and geometry are similar to the PAM ones, the
fabrication process is similar. The differences are only the transducer dimensions and
focusing configuration. For endoscopic applications, the PAE transducer was connected to
a core wire of 0.44 mm thick, 120 cm long, 50 Ω coaxial cable (its shield wires are
connected to the housing). The acoustic focusing was achieved by attaching a plano-
concave plastic acoustic lens to the flat surface of the US transducer. The plastic acoustic
lens was fabricated in-house by molding polyester resin. Figure 5-5 shows photos of the
implemented US transducer in which the piezo-element has an outer diameter of ~2.6 mm
and an inner diameter of ~0.7 mm.
-2
-1
0
1
2
-2
-1
0
1
2
0
0.5
1

x [mm]
y [mm]

z [mm]
-40
-40
-33
-33
-33
-33
-33
-33
-33
-33
-27
-27
-27
-27
-27
-27
-27
-27
-27
-27
-27
-27
-27
-21
-21
-21
-21
-21
-21
-21
-21
-21
-21
-21
-21
-21
-21
-15
-15
-15
-15
-15
-12
-12
-12
-12
-9
-9
-6
-3
-3
Distance (mm) Along Beam Axis
Distance (mm) Perpendicular to Beam Axis
-1.5 -1 -0.5 0 0.5 1 1.5
0
2
4
6
8
10
12
54

Figure 5-5: Photos of a finished 1
st
generation PAE transducer.

5.2.2 Transducer Performance Evaluation
Measured electrical impedance and phase of the 1
st
generation PAE transducer are
shown in Figure 5-6. The electrical impedance at a phase peak was found to be 133.6 Ω at
47.9 MHz. The series and parallel resonant frequencies are 43.4 MHz and 53.0 MHz,
respectively. The values are comparable to the modeling prediction. Measured pulse-echo
waveform and frequency spectrum of the transducer are shown in Figure 5-7. The center
frequency of the transducer was 40 MHz and the -6 dB bandwidth was 80 %. Compensated
insertion loss of the transducer was measured to be 17.3 dB. The experimental results show
that the transducer made for PAE application is much more sensitive than the commercial
Panametrics single element transducer (Olympus, Waltham, MA, USA). By placing optical
components in the hole of the ring transducer, axially-symmetric common axis of optical
illumination and ultrasound detection can be obtained.
55

Figure 5-6: Measured electrical impedance and phase of the 1
st
generation PAE
transducer.

Figure 5-7: Measured pulse-echo impulse reponse and its FFT spectrum of the 1
st

generation PAE transducer.
5.2.3 Ultrasound Resolution PAE Probe Implementation
Figure 5-8 shows (a) a schematic and (b) a photograph of the distal end of the
photoacoustic endoscopic probe. In the probe, the PAE transducer, a light guiding optical
20 30 40 50 60 70 80
100
125
150
175
200
Frequency (MHz)
Electrical Impedance ( )

20 30 40 50 60 70 80
-90
-72.5
-55
-37.5
-20
Phase (deg)
Electrical Impedance
Phase
8.0645 8.1898 8.315 8.4403 8.5655
-400
-200
0
200
400
Time ( s)
Amplitude (mV)

10 25 40 55 70
-24
-18
-12
-6
0
Frequency (MHz)
Magnitude (dB)
Pulse-echo Response
Spectrum
56

fiber (0.22 NA, 365 µ m core dia.), and a mechanical micro-motor were placed into a
stainless steel tube. Circumferential sector scanning (B-scan) was accomplished by rotating
a mirror (3.0 mm diameter, protected aluminum on glass substrate with the reflection
surface at 45° to the probe’s axis). The mirror - driven by a 1.5 mm diameter, 12.0 mm
long geared micro-motor (gear ratio, 254:1; Namiki Precision Inc.) – was used to steer both
the light beam from the optical fiber to the tissue and the photoacoustic wave from the
tissue to the transducer. The optical fiber, the transducer’s signal wires, and the micro-
motor’s wires were encapsulated in the flexible endoscope body. The outer diameter of the
implemented probe was 3.8 mm. One of the important features of the imaging probe was
its built-in scanning mirror-based actuator housed at the distal end of the probe, which
enabled static mounting of the associated illumination and ultrasound pulse generation-
detection units. This built-in scanning mechanism yields higher-quality images because
mechanical scanning is much more stable than the conventional, flexible-shaft-based,
proximal-actuation method.

Figure 5-8: (a) Schematic of the photoacoustic endoscope’s distal end. GM, geared
micromotor; JB, jewel bearings; MN, magnets; OF, optical fiber; PM, plastic membrane
(imaging window); SM, scanning mirror; UST, ultrasonic transducer. (b) Photograph of
the probe. (c) Field of view. SW, stainless steel wall (blocked zone: 110° ); PM, plastic
membrane (imaging zone: 250° ).
57

5.2.4 In vivo PAE Imaging
The US images are produced with conventional pulse-echo imaging that detects
acoustic waves reflected from target tissue; the PA images are formed through detection of
acoustic waves generated by rapid thermoelastic expansion caused by optical absorption
of short laser pulses. The focused US transducer detected one-dimensional depth-resolved
signals (or A-lines), and cross-sectional images (or B-scans) were produced by rotating a
scanning mirror which directs both optical and acoustic waves. The endoscope system
recorded and displayed a set of dual wavelength PA and US B-scan images in real-time
during the constant rotation of the mirror at frequency of about 4 Hz. By interleaving two
optical pulses of different wavelengths and one acoustic pulse at each angular step of the
mirror, spatially-coincident images were recorded from the generated PA and US A-line
signals, even during periods of significant motion of the target. Volumetric data sets were
acquired by recording sequential A-line data during the constant rotational motion of the
mirror and mechanical pullback of the probe at a speed of ~200 µ m s
-1
.
Figure 5-9 illustrates of simultaneous, multi-wavelength photoacoustic (PA) and
ultrasonic (US) endoscopy. The endoscope performed circumferential sector scanning by
rotating a scanning mirror, which reflected both the US waves and laser pulses and enables
static mounting of the associated illumination and US pulse generation-detection units
(Figure 5-9(a)). At each angular step of the mirror (~1.42°), both the first (λ1) and second
(λ2) pulsed laser beams were independently fired through the optical fiber and the acoustic
pulse was generated by the US transducer with a constant time delay of ~30 µ s between
each of the laser and acoustic pulses. The ensuing PA and US echo waves were detected
and converted into electric signals by the US transducer; the signals were then recorded
58

and displayed on a computer. Figure 5-9(b) shows the side-scanning 3.8-mm diameter
probe prototype firing a 562 nm laser beam. Figure 5-9(c) is the definition of Cartesian and
cylindrical coordinate systems. The +z-axis is defined along the endoscope axis (or
pullback direction). Figure 5-9(d) shows that a volumetric image comprised of consecutive
B-scan slices. Figure 5-9(e) shows a representative cross-section of Figure 5-9(d) along the
x-y plane, which shows 270° angular field of view of the endoscope.

Figure 5-9: Illustration of simultaneous, multi-wavelength photoacoustic (PA) and
ultrasonic (US) endoscopy.

The endoscopic potential of the probe was proven by imaging the esophageal tract
of an adult rabbit (New Zealand white, 4.0 kg) in vivo. During the in vivo imaging
procedure, anesthesia gas (1.5-2.0% of isoflurane) and oxygen were continuously supplied
through an intubation catheter. The catheter had a balloon positioned near the branching
59

point between the trachea and esophageal tract, which avoided water invasion to the lung
as the esophageal tract was filled with water for acoustic matching. The endoscopic probe
was inserted into the esophageal tract ~25 cm deep from the animal’s mouth and performed
pullback C-scans over a ~15 cm range during the scanning time of ~10 min for each. About
2800 B-scan slices with a longitudinal spacing of ~50 µ m were acquired for each data set
of PA (584 nm), and US to produce volumetric images. For the PA and US imaging, the
radial imaging depth was ~7 mm from the surface of the endoscope (i.e., 18 mm diameter
FOV). Since the angular FOV of the endoscope was partially restricted, the endoscope
direction was adjusted appropriately to target mainly the ventral side of the animal.
Experimentally measured PA and US resolutions in the focal zone of the transducer were
respectively ~55 µ m and ~30 µ m in the radial direction, and ~80 µ m and ~60 µ m in the
transverse direction.
Figure 5-10 shows the PA (Figure 5-10(a)) and US (Figure 5-10(b)) images, and a
merged image (Figure 5-10(c)) from a volume of 15 cm long and 18 mm diameter in a
rabbit esophageal tract. In both the PA and US images, not only the esophageal tract, but
also other surrounding organs, such as the lung (left arrow) and trachea tract (right arrow),
which contacted with the esophageal tract, was imaged clearly. However, only the PA
image clearly shows the blood vasculature surrounding these organs. The combined image
(Figure 5-10(c)) shows the contrast difference between the PA and US images clearly. The
positions and sizes of the organs shown in the images coincided well with the actual
anatomical distributions confirmed surgically after the imaging experiment.
60

Figure 5-10: Simultaneous, co-registered PA and US endoscopy of a rabbit esophageal
tract in vivo over an 18 mm diameter and 15 cm long FOV. The lower parts of all images
correspond to the ventral part of the animal. Horizontal and vertical scale bars, 2 cm and 5
mm, respectively. (a) Three-dimensionally-rendered PA image acquired at 584-nm
(isosbestic point) wavelength, showing the total hemoglobin distribution. (b) Three-
dimensionally-rendered, co-registered US structural image for the same volume of (a). (c)
A combined image of (a) and (b).
5.3 Second Generation Transducer Design: Self-focused Ring Transducer
5.3.1 Transducer Design and Fabrication Process
Since the laser light in the PAE probe based on the first generation PAE transducer
is not focused, the imaging lateral resolution can be achieved only acoustic-resolution as
described in Chapter 2. To achieve optical resolution PAE (OR-PAE) which can provide
greatly improved lateral resolution, both the ultrasonic transducer and laser light beam should
be focused coaxially. Furthermore, to avoid acoustic attenuation caused by the lens, another
transducer focusing technique should be utilized without lens. Therefore, a self-focus
transducer that can also focus the laser light beam has been developed. The schematic of
the transducer is shown in Figure 5-11. The fabrication process was similar to the first
generation PAE transducer. However, the back-end of the transducer was drilled with a
hole with 1.3 mm diameter for providing enough space to place a tiny homemade optical
lens inside the transducer. Moreover, surface of the transducer was focused by hot pressing
61

a ball bearing of radius equal to the desired imaging focal distance which was 5 mm. A
finished second generation PAE transducer is shown in Figure 5-12.

Figure 5-11: Schematic of 2
nd
generation PAE transducer.

Figure 5-12: Photos of a finished 2
nd
generation PAE transducer.

5.3.2 Transducer Performance Evaluation
Measured pulse-echo waveform and frequency spectrum of the second generation
transducer are shown in Figure 5-13. The center frequency of the transducer was measured
62

at 38.5 MHz and the -6 dB bandwidth was 90 %. Compensated insertion loss of the
transducer was measured to be 18.7 dB.

Figure 5-13: Measured pulse-echo impulse reponse and its FFT spectrum of the 2
nd

generation PAE transducer.

5.4 Third Generation Transducer Design: Self-focused Circular Shape Transducer
5.4.1 Transducer Design and Fabrication Process
As shown above, 3.8-mm diameter integrated photoacoustic and ultrasonic
endoscopic probe was demonstrated. However, its rigid distal sections are too long to be
inserted in the instrument channel (typically ~2.8 mm or ~3.7 mm in diameter) of a standard
video endoscope because they work based on a micromotor for mechanical scanning of
6.8395 6.9648 7.09 7.2153 7.3405
-400
-200
0
200
400
Time ( s)
Amplitude (mV)

10 25 40 55 70
-24
-18
-12
-6
0
Frequency (MHz)
Magnitude (dB)
Pulse-echo Response
Spectrum
63

mirror in front of the transducer. Thus, developing new endoscopic probes with improved
flexibility and imaging capabilities is necessary. Hence, a self-focused circular shape
transducer with smaller aperture size 2.0 mm was developed and employed in a flexible
shaft-based PAE probe. The fabrication process was similar to the first and second
generation PAE transducers. However, inner hole is not needed for flexible shaft-based
PAE probe. Therefore, this transducer has more uniform acoustic field than first and second
generation PAE transducers regardless of effect induced by inner hole. A finish third
generation PAE transducer is shown in Figure 5-14.

Figure 5-14: Photos of a finished 3
rd
generation PAE transducer.

5.4.2 Transducer Performance Evaluation
Measured pulse-echo waveform and frequency spectrum of the third generation
transducer are shown in Figure 5-15. The center frequency of the transducer was measured
at 51.0 MHz and the -6 dB bandwidth was 90 %. Compensated insertion loss of the
transducer was measured to be 32.4 dB. Since the total area of the transducer was reduced,
64

the sensitivity of the third generation PAE transducer is less than the first and second PAE
transducers because of higher electrical impedance mismatch.

Figure 5-15: Measured pulse-echo impulse reponse and its FFT spectrum of the 3
rd

generation PAE transducer.

5.4.3 Flexible Shaft-based PAE Probe Implementation
As shown in the Figure 5-16, the flexible shaft-based PAE endoscopic system is
comprised of a proximal actuation unit, a ~2.5-m long flexible body section sheathed with
a plastic tube, and a rigid distal section that has a rotatable scanning mirror. In this
mechanical scanning endoscopic system, a step motor installed in the actuation unit serves
as a mechanical source that providing a torque. The torque is transferred to a metal shaft
sustained by two ball bearings through a mechanical engagement (timing pulley and belt),
5.2395 5.3648 5.49 5.6152 5.7405
-200
-100
0
100
200
Time ( s)
Amplitude (mV)

0 25 50 75 100
-24
-18
-12
-6
0
Frequency (MHz)
Magnitude (dB)
Pulse-echo Response
Spectrum
65

and further transmitted to the scanning mirror through a ~2.5-m long flexible shaft. Finally,
the scanning mirror performs a rotational scanning (~5 Hz) receiving the torque.
For PA imaging, laser pulses from a laser source (Nd:YVO4, INNOSLAB IS811-
E, EdgeWave) are guided by a multimode optical fiber (0.22 NA, 365 µ m core diameter,
BFL22-365, Thorlabs), transferred to the endoscope’s optical fiber located inside the
flexible shaft via a disjointed rotary junction (coupling), and finally sent to target tissue by
the scanning mirror. Some of the PA waves that propagate to the scanning mirror are
reflected by the same mirror, sent to the transducer, and converted into electrical signals.

Figure 5-16: A diagram of the flexible shaft-based PAE.

5.5 Summary
In this chapter, three types of LiNbO3 ultrasonic transducers including lens-focused
ring shape, self-focused ring shape, and self-focused circular shape have been designed,
fabricated, and characterized for PAE application. The lens-focused ring shape transducer
66

provides axially-symmetric common axis of optical illumination and ultrasound detection
by placing optical components at the inner hole of transducer. The imaging feasibility of
the PAE probe based on lens-focused ring shape transducer has been demonstrated through
in vivo volumetric imaging of rabbit esophageal tract (Yang et al. 2012). The self-focused
ring shape transducer, capable of focusing laser light beam by inner installed optical lens,
can be used to obtain optical-resolution PA imaging. Small aperture self-focused circular
shape transducer was employed in a shaft-based PAE probe to improved flexibility and
imaging capabilities by being inserted in the instrument channel of a standard video
endoscope.

67

Chapter 6 High-frequency Phased Array for Photoacoustic
Endoscopy Application
6.1 Introduction
As described in Chapter 5, single-element transducer based side-looking (SL)
probes provide high-resolution cross-sectional images. However, motion artifacts may be
induced to affect image quality because of mechanical scanning of the single-element
transducers, and the image quality is best only at the transducer focus. SL probes may also
use circular arrays mounted along the circumference of the catheter to enable electronic
scanning of the cross-sectional area without mechanical artifact caused by rotating
transducers (O’Donnell et al. 1997). A major disadvantage of these SL probes is the lack
of forward-looking (FL) capability to provide a view of the path or structures in front of the
catheter that are helpful in guiding interventions, especially in the case of chronic
total occlusions (Tekes et al. 2007). Although single-element transducer based systems in
which the transducer is mounted on a rotating cam assembly have been realized for FL,
these systems require a complex mechanism with complicated beam alignment and
calibration protocols (Evans et al. 1994; Ng et al. 1994; Liang and Hu1997). Therefore, the
development of a miniature high-frequency array transducer may be beneficial in FL-PAE
imaging systems for providing more diagnostic capabilities and device maneuvering.
Further, phased array is more suitable than linear array for FL-PAE applications because it
is capable of providing a wider field of view in the far field.
68

Although in recent years, there has been intensive development in high-frequency
linear arrays, the physical limitation in fabrication technology restricts the development of
high-frequency arrays, especially for intravascular imaging applications. In conventional
piezoelectric transducer technology, the arrays are prepared from mechanical dicing or
laser-dicing of a plate of transducer material so that the elements are separated physically
(kerfed arrays). Previously, mechanically diced linear arrays with center frequencies up to
30 MHz have been reported (Nguyen-Dinh et al. 1996; Michau et al. 2004). However, the
pitch of these arrays was greater than 1/2  which was too large to satisfy the criteria of
phased array design. By adopting laser-dicing technique, the kerf can be minimized so that
PZT arrays with center frequencies up to 50 MHz have been developed successfully by
Foster’s group for small animal imaging (Foster et al. 2000; Foster et al. 2009; Lukacs et
al. 2006). However, the pitch is still too large and the entire array size is too big for PAE
imaging applications. More recently, a 64-element 35-MHz composite ultrasonic array was
developed using a mechanical “dice-and-fill” method (Cannata et al. 2006). The pitch of
the array was successfully pushed to the almost 1/2  or 25 μm using the double index
dicing technique. In order to construct miniature high-frequency array transducers,
capacitive micromachined ultrasonic transducers (CMUT) were also studied extensively.
Since the array size can be minimized easily with integrated circuit (IC) technology with
CMUT, researchers have put forward great effort in developing forward-looking circular
arrays for intravascular imaging (Tekes et al. 2011; Oralkan et al. 2004; Degertekin et al.
2006; Yeh et al. 2006). These devices have very complex configuration that utilizes sparse
array approach. Thus far the highest frequency reported is only at 20 MHz.
69

As an alternative to mechanical separation of transducer elements physically,
ultrasonic arrays can also be constructed in separating the elements electronically which
are so-called kerfless arrays (Morton and Lockwood 2002). The electrodes of the array
elements are simply patterned onto the transducer surface. Initially, the kerfless arrays were
developed for annular arrays because the circular pattern cannot be made with a dicing saw
(Morton and Lockwood 2001). Compared to the kerfed arrays, the fabrication method of
the kerfless arrays is simple and reliable but the drawbacks are relatively high cross-talk
and large active area. Nevertheless, the kerfless arrays have been shown to exhibit
comparable performance to the kerfed arrays if the undesired factors are considered in the
system design (Harman 2010). Recently, Bezanson et al. developed a 40-MHz phased array
using a FL kerfless design (Bezanson et al. 2012). Although the transducer performance
has been improved significantly compared to the previous reports, the pitch is larger than
1/2  and the array size is large.
Besides the fabrication technology, materials also play an important role on the
array transducer performance. Previously, piezoelectric lead zirconate titanate (abbreviated
as PZT) ceramics and films were used for developing high-frequency kerfless linear arrays
(Wu et al. 2009). Although the array frequency can be increased easily using the thick film
as the transducer element, the insertion loss is generally higher because of the poor material
quality of the film. Bulk ceramics are much denser that films so as to exhibit much better
performance. The transducer performance was reported to be enhanced when bulk
ceramics was the transducer piezoelectric element. Besides piezoelectric ceramics, lead
magnesium niobate-lead titanate (PMN-PT) single crystals near the morphotropic phase
boundary (MPB) composition have also been used for wide range of applications (Park and
70

Shrout 1997; Lam et al. 2005; Edwards et al. 2006). The features of high piezoelectric
capability, high dielectric constant and low dielectric loss make the PMN-PT single crystals
ideal for high-sensitive and small-aperture transducers design (Lau et al. 2004; Zhou et al.
2007; Lam et al. 2012). However, PMN-PT single crystals are more brittle in nature and
prone to fracture than PZT ceramics. It may cause excessive cracking during mechanical
dicing process for kerfed array especially with a small pitch.
The aim of this work is to study the development of a miniature 32-element high-
frequency kerfless phased arrays using PMN-PT single crystal and a miniature 32-element
high-frequency kerfled phased arrays using PZT ceramics. The design, fabrication, and
characterization of both arrays are presented in this paper. The axial and lateral resolutions
of the kerfless array were evaluated by ultrasound (US) imaging of a wire phantom. Its
photoacoustic (PA) imaging capability has been demonstrated by imaging a graphite rod
target.
6.2 High-frequency PMN-PT Kerfless Phased Array
6.2.1 Design
To achieve adequate signal sensitivity with the restricted element size, a highly
sensitive PMN-PT single crystal was selected as piezoelectric material, which has high
dielectric permittivity (
33 0
/
s
 ) and high electromechanical coupling factor (
t
k ). Table 6-1
lists major properties of the PMN-PT single crystal used in this study. The array transducer
was targeted at a center frequency of around 40 MHz for high resolution capability. The
71

design parameters of the kerfless array are shown in Table 6-2. Figure 6-1 shows the
schematic of the phased array transducer.

Table 6-1: Properties of PMN-PT single crystal.
Material PMN-30%PT single crystal
Vendor H. C. Materials Corp.
Density 7800 kg/m
3

Acoustic Impedance 37 MRayl
Piezoelectric d 33 coefficient 1500 pC/N
Relative Dielectric Constant (free) ~ 5000
Relative Dielectric Constant (clamped) 800
Loss Tangent 0.005
Electromechanical Coupling Coefficient k t 0.58
Acoustic velocity 4600 m/s

Figure 6-1: Schematic of the phased array transducer (L: length of array, H: height of
array, W: element width, K: kerf width, P: pitch).
72

Table 6-2: Design parameters of the phased array transducer.
Specifications Values
Design Frequency 40 MHz
Pillar Width 25 μm
Kerf 8 μm
Pitch 33 μm
Height 1 mm
Length 0.8 mm
Number of Elements 32
Piezoelectric Material (PM) PMN-PT single crystal
Thickness of PM 40 μm
Matching Layer (ML) parylene
Thickness of ML 12 μm
Backing E-solder 3022
Thickness of Backing 2 mm

To validate the array design, a 2-D finite element model (PZFlex, Weidlinger
Associates Inc., Mountain View, CA) was used to predict the electrical impedance, pulse-echo
response, and cross-talk performance of the array. The finite element modeling provides an
accurate prediction of high-frequency array performance, as well as reduce the number of
time-consuming prototype fabrication runs.
6.2.2 Fabrication Process
The high-frequency kerfless phased array transducer was fabricated using
micromachining technique. Specifically, a <001> oriented bulk PMN-33%PT single
crystal (H. C. Materials Corporation, Bolingbrook, IL, USA) was polished from one side.
A thin layer of Cr/Au (500Å/1000Å) film was sputtered on the polished side as top
electrodes. Here, a photo-mask was designed for patterning the 32-element phased arrays.
Figure 6-2 shows the layout of the mask, where the phased array has a kerf of 8 μm, an
73

element width of 25 μm and an element length of 1 mm. The design patterns were then
patterned onto the surface using a combination of photolithography and Cr/Au etching.
After patterning top electrodes, the samples were flipped over to the other side and lapped
to the designed thickness of ~40 μm. A 2 mm thick conductive backing material, E-solder
3022 (VonRoll Isola, New Haven, CT, USA), was cured on the other side of the PMN-PT
single crystal. These processing procedures are illustrated in Figure 6-3. Each array
element was then individually connected to a thin wire by a small amount of E-Solder. The
arrays were housed in an aluminum tube and sealed with insulated epoxy (EPO-TEK 301,
Epoxy Technology, Inc., Billerica, MA). Finally, vapor-deposited parylene with a
thickness of 12 μm was used to coat the aperture and the housing. Figure 6-4 shows a photo
of the kerfless phase array transducer prototype. The active area of the array is 0.8 mm ×
1.0 mm. After fabrication, every array elements was poled in air at room temperature under
an electric field of 20 kV/cm for 10 min.

Figure 6-2: The array pattern with 32-elements.
74

Figure 6-3: Fabrication flow of PMN-PT kerfless phased array.

Figure 6-4: Photo of the PMN-PT kerfless phase array prototype.

75

6.2.3 Performance Evaluation
Several standard non-imaging transducer tests were first performed on the array to
characterize its performance (Cannata et al. 2006).
The simulated and measured frequency dependence of the electrical impedance and
phase of a representative array element are displayed in Figure 6-5 and Figure 6-6,
respectively. The simulated and measured results are compared in Table 6-3. The measured
results showed that the electrical impedance at a phase peak was 69.9 ± 2.9 Ω at 40 MHz.
The series and parallel resonant frequencies were 39.4 ± 1.0 MHz and 46.1 ± 1.1 MHz,
respectively. The electromechanical coupling coefficient, kt of the array elements was
found to be 0.56 ± 0.02 which was comparable to that of the bulk PMN-PT single crystal
(0.58, HC materials, Bolingbrook, IL, USA).

Figure 6-5: Simulated electrical impedance and phase of a representative kerfless array
element by PZFlex.
76

Figure 6-6: Measured electrical impedance and phase of a representative kerfless array
element.

Table 6-3: Comparison of simulated and measured electrical impedance results.
Property PZFlex Measured
Impedance 85.9 Ω @ 40 MHz 69.9 ± 2.9 Ω @ 40 MHz
f r 37.3 MHz 39.4 ± 1.0 MHz
f a 45.5 MHz 45.7 ± 1.1 MHz
k t 0.61 0.56 ± 0.02

The simulated and measured pulse-echo responses and their corresponding FFT
spectra of a representative array element are shown in Figure 6-7 and Figure 6-8,
respectively. The center frequency and -6 dB bandwidth of 32 array elements was
measured at 42.6 ± 0.4 MHz and 34.1 ± 2.2%, respectively. The measured insertion loss
for all array elements was 20.0 ± 1.3 dB. Comparison of simulated and measured results is
shown in Table 6-4.
77

Figure 6-7: Simulated pulse-echo response and its FFT spectrum of a representative
kerfless array element.

Figure 6-8: Measured pulse-echo response and its FFT spectrum of a representative
kerfless array element.

3.5945 3.8448 4.095 4.3453 4.5955
-500
-250
0
250
500
Time ( s)
Amplitude (mV)

20 30 40 50 60
-24
-18
-12
-6
0
Frequency (MHz)
Magnitude (dB)
Pulse-echo Response
Spectrum
78

Table 6-4: Comparison of simulated and measured pulse-echo response and its FFT
spectrum.
Property PZFlex Measured
f
c
(MHz)
41.5 ± 0.6 42.6 ± 0.4
-6 dB BW (%) 31.3 ± 1.3 34.1 ± 2.2
V pp (mV) 534.8 ± 127.5 653.7 ± 100.3
-6 dB / -20 dB pulse length 69 / 147 nsec 63 / 169 nsec
Compensated IL (dB) N/A 20.0 ± 1.3

The simulated and measured crosstalk for the array are shown in Figure 6-9(a) and
6-9(b), respectively. The measured crosstalk near the center frequency of the array was <
−10 dB which is higher than that of the conventional kerfed arrays (Cannata et al. 2005).

Figure 6-9: PZFlex simulated (a) and measured (b) cross-talk of the array.

Compared to the conventional kerfed arrays, the cross-talk of the kerfless array is
high so as to result in a reduction of the element acceptance angle. The measured one-way
directivity pattern for a representative array element is shown in Figure 6-10. -6 dB
acceptance angle was 10º for the kerfless array element. As expected, the kerfless array
directivity suffered from increased crosstalk between elements.
79

Figure 6-10: Measured one-way directivity for a single array element of the kerfless
array.

The measured results of the kerfless array are summarized and compared with
PZFlex modeling results in Table 6-5. It is obvious that the measured results are in
accordance to the measured results.

Table 6-5: Comparison between measured and modeling results for the array.
Property Measured PZFlex
# of elements 32 32
# of open/shorted elements 0/0 0/0
Average f c 42.6 MHz 41.3 MHz
Highest/lowest f c 43.4 MHz / 41.7 MHz 42.1 MHz / 40.4 MHz
Average -6dB BW 34 % 29 %
Highest/lowest -6dB BW 37.3 % / 29.0 % 32.1 % / 27.5 %
Average V p-p 653.7 mV (no gain) 534.8 mV (no gain)
Highest/lowest V p-p 898.0 mV / 391.5 mV 654.2 mV / 162.6 mV
Average IL 20.0 dB N/A
Highest/lowest IL 23.7 dB / 18.1 dB N/A
-6 dB / -20 dB pulse length 63 / 169 nsec 69 / 147 nsec
k t 0.557 0.612
Crosstalk < -5 dB < -5 dB
80

6.2.4 Imaging Test
The ultimate performance indication of a transducer is determined by its imaging
capability. In the imaging evaluation, the kerfless phased array was paired with a 32-
channel high-frequency ultrasound imaging system to image a fine wire phantom. This
approach was used to determine the array’s spatial resolution.
The customized digital imaging system sampled the echo signals at 140 mega samples
per second (MSPS). Totally 128 scanning beams were focused and steered within ± 30º .
No apodization or thresholding was implemented during imaging reconstruction. Some
specifications of the imaging system are given in
Table 6-6.

Table 6-6: Specifications of 32-channel digital imaging system.
Number of transmit channels 32
Number of receive channels 32
Number of bits per channel 12
Number of samples per scan line 2048 (around 11 mm depth)
Transmit focus one focus at 4mm
Receive focus Dynamic, updated with every sample

The imaging target was five 20μm diameter tungsten wires, (California Fine Wire
Company, Grover Beach, CA) which were arranged diagonally with equal distance in the
axial (1.5 mm) and lateral (0.65 mm) directions, respectively (Figure 6-11). In the
experiment, the wire phantom target was immersed in degassed water.
81

Figure 6-11: The arrangement of five tungsten wire targets.

The B-mode image of this five-wire phantom is shown in Figure 7, which is in a
linear gray scale and 50-dB dynamic range. Plots of the axial and lateral line spread
functions for the 2nd wire are shown in Figure 6-13. The measured full-width half-
maximum (FWHM) spatial resolutions were 467 µ m and 118 µ m in lateral and axial
directions, respectively. For comparison, Field II (Jensen 1992; Jensen 1996) simulation is
shown in Figure 6-14. The lateral and axial profiles of the second wire are shown in Figure
6-15. The theoretical lateral and axial resolution are 229 µ m and 90 µ m, respectively. Large
discrepancies between theoretical (Field II) and measured resolutions are expected due to
the presence of high cross-talk in kerfless array. Furthermore, wires were simulated as
perfect point sources and the array itself was considered as kerfed array in Field II.
Tungsten wires
(20 um diameter)
De-ionized
water
.65
mm
1.5
mm
1.5
mm
1.5
mm
1.5
mm
.65
mm
.65
mm
.65
mm
82

Figure 6-12: Wire phantom imaging with the 40 MHz PMN-PT kerfless phased array.

Figure 6-13: Lateral (left) and axial (right) line spread functions for the 2
nd
wire of the
wire phantom.

Depth (mm)
Lateral distance (mm)

-5 0 5
1
2
3
4
5
6
7
8
9
10
11
5
10
15
20
25
30
35
40
45
50
83

Figure 6-14: Field II simulated wire phantom image with 40 MHz phased array.

Figure 6-15: Field II simulated lateral (left) and axial (right) line spread functions for the
2nd wire of the wire phantom.

This kerfless phased array was also evaluated for photoacoustic (PA) imaging since
photoacoustic signals are subject to only one-way, the impact to the imaging quality caused
by cross-talk is expected to be less than round-trip. A graphite rod with 0.5 mm diameter
Lateral distance [mm]
Depth [mm]

-5 0 5
1
2
3
4
5
6
7
8
9
10
11
5
10
15
20
25
30
35
40
45
50
84

was used as a PA imaging target which has strong light absorption characteristics (Kong et
al. 2009). In our PA imaging system, a pulsed Q-switched Nd:YAG laser (Spectra-Physics
Explorer 532-2Y, Newport Corp., Santa Clara, CA) was used as PA excitation source. The
free space laser output was coupled by a 4X objective lens into an optical fiber and then
delivered to the lead strip surface. The array was placed in front of the graphite rod and
connected to the high-frequency ultrasonic phased array system. The high-frequency
ultrasonic phased array system was used to generate US pulses and receiving both US and
PA signals. Received signals are digitized and processed in the computer. A co-registered
US and PA image of the graphite rod is shown in Figure 6-16 using a linear gray scale and
50-dB dynamic range. The scan angle is also set at ± 30
o
. No apodization or thresholding
was implemented during reconstruction.

Figure 6-16: Co-registered US and PA image of a graphite rod using the 40-MHz PMN-
PT single crystal kerfless phased array.

85

6.3 High-frequency PZT Kerfed Phased Array
6.3.1 Design
Although kerfless phased array has been demonstrated potential for PAE
application as presented above, the drawbacks are obvious including high crosstalk and
small acceptance angle. Thus, development of kerfed array for PAE application is
necessary because of its better overall performance than kerfless array.
A 32-elements high-frequency phased array using 2-2 piezo-composite has been
designed and developed. Design parameters of the kerfed array are listed in Table 6-7.
Dice-and-fill (Cannata et al. 2006) and interdigitally bonding techniques (Cannata et al.
2011) were utilized to prepare 2-2 piezo-composite to overcome kerf width limitation of
the mechanical dicing saw. Although the features of high piezoelectric capability, high
dielectric constant and low dielectric loss make the PMN-PT single crystals are ideal for
high-sensitive and small-aperture transducers design, PMN-PT single crystals are more
brittle in nature with fracture toughness than PZT ceramics. It may cause excessive
cracking during mechanical dicing process for kerfed array especially with a small pitch.
Thus, PZT-5H ceramic was chosen as the piezoelectric material for the kerfed phased array.
Figure 6-17 shows the prepared PZT-5H (TFT-L201F, TFT Corporation,Tokyo, Japan) 2-
2 composite with 25 μm pitch which equals to half-wavelength at 30 MHz and 6 μm kerf.
Half-wavelength element pitch prevents the generation of grating lobes.

86

Table 6-7: Design parameters of the kerfed phased array transducer.
Specifications Values
Design Frequency 30 MHz
Element Width 19 μm
Kerf 6 μm
Pitch 25 μm
Height 1 mm
Length 0.8 mm
Number of Elements 32
Piezoelectric Material (PM) PZT 5H
Thickness of PM 55 μm
Matching Layer (ML) parylene
Thickness of ML 17 μm
Backing E-solder 3022
Thickness of Backing 2 mm

Figure 6-17: PZT-5H 2-2 composite.

6.3.2 Fabrication Process
A flex-circuit (Figure 6-18) was fabricated at cleanroom of photonics center at USC
with a 25 μm thick kapton film. The electrode on the flex-circuit was achieved using deep
reactive-ion etching (DRIE) method. A photo-mask was designed for patterning the
87

electrode. The narrow trace width at the center was 10 μm. After alignment the flex circuit
was bonded in front of the array surface with insulating epoxy (Epo-Tek 301, Epoxy
Technologies,Billerica, MA) and bended over and inserted into a tube shape small stainless
steel housing. The outer diameter of the housing was about 2 mm. The front surface of
finished kerfed array is shown in Figure 6-19.

Figure 6-18: Flex-circuit for the kerfed array.

Figure 6-19: The front surface of the kerfed array.

88

6.3.3 Performance Evaluation
The measured frequency dependence of the electrical impedance and phase of a
representative kerfed array element is shown in Figure 6-20. The measured results showed
that the electrical impedance at a phase peak was 898.4 Ω at 30 MHz. The series and
parallel resonant frequencies were 29.2 MHz and 34.0 MHz, respectively. kt of the array
was found to be 0.55.

Figure 6-20: Measured electrical impedance and phase of a representative kerfed array
element.

The measured pulse-echo responses and their corresponding FFT spectrum of a
representative kerfed array element are shown in Figure 6-21. The average center
frequency and -6 dB bandwidth of 32 array elements was measured to be 32.5 MHz and
52 %, respectively. The average measured insertion loss for all array elements was 60.6
dB.
10 15 20 25 30 35 40 45 50
500
1000
1500
2000
2500
Frequency (MHz)
Electrical Impedance ( )

10 15 20 25 30 35 40 45 50
-90
-83.75
-77.5
-71.25
-65
Phase (deg)
Electrical Impedance
Phase
89

Figure 6-21: Measured pulse-echo response and its FFT spectrum of a representative
kerfled array element.
6.4 Summary
A 40-MHz 32-element kerfless phased array using PMN-PT single crystal and a
30-MHz 32-element kerfed phased array using PZT-5H ceramic were modeled, fabricated
and tested. Kerfless array’s imaging capability was demonstrated with a US wire phantom
and PA a graphite rod phantom. By using the PMN-PT single crystal as the transducer
element, this kerfless array is much sensitive than the one previously built (Cannata et al.
2005). Their performance are summarized and shown in Table 6-8. As expected, the
kerfless array directivity suffered from the high crosstalk between elements. These results
suggested that kerfless array technology may be a viable alternative for high-frequency
phased array designs. The miniature active area of both arrays makes them good candidates
for PAE application.
1.6495 1.8998 2.15 2.4002 2.6505
-100
-50
0
50
100
Time ( s)
Amplitude (mV)

10 20 30 40 50
-24
-18
-12
-6
0
Frequency (MHz)
Magnitude (dB)
Pulse-echo Response
Spectrum
90

Table 6-8: Phased array performance summary.
Property
Kerfless
Array
Kerfed Array
Previous Built Kerfless
Array (Cannata 2005)
Material
PMN-PT
single crystal
PZT ceramic 2-2
composite
PZT ceramic
Number of elements 32 32 64
Number of
open/shorted elements
0/0 8/0 13/2
Average center
frequency
42.6 MHz 32.5 MHz 32.0 MHz
Average bandwidth (-
6dB)
34 % 49 % 61 %
Average sensitivity
653.7 mV (no
gain)
101.1 mV (20 dB
gain)
1.65 V (20 dB gain)
Average insertion loss 20.0 dB 61.5 dB 19.9 dB
-6 dB / -20 dB pulse
length
63 / 169 nsec 43 / 107 nsec 40 / 124 nsec
Cross-talk < -5 dB N/A < -11 dB
-6 dB transmit
acceptance angle
12º N/A 22º

91

Chapter 7 Summary and Future Work
7.1 Summary
In this dissertation, efforts in developing novel high-frequency ultrasonic
transducers for PA imaging modalities are reported. The results demonstrate that it is
possible to improve spatial resolution with these designs.
For PAM application, a highly focus LiNbO3 ring-shape ultrasonic transducer with
2.4 mm inner diameter and 4.8 mm outer diameter has been designed and fabricated for
UV-PAM application. From the performance evaluation, the center frequency, -6 dB
bandwidth, and insertion loss of the transducer have been measured. The results suggest
that this transducer is very suitable for high resolution PAM imaging application. The ex
vivo PAM image of micron-level cell nuclei acquired from this transducer demonstrated
the significant improved lateral resolution which achieve submicron level with specific,
positive, and high contrast.
For PAE application, three types of LiNbO3 ultrasonic transducers including lens-
focused ring shape, self-focused ring shape, and self-focused circular shape have been
designed, fabricated, and characterized for PAE application. The lens-focused ring shape
transducer provides axially-symmetric common axis of optical illumination and ultrasound
detection by placing optical components at the inner hole of transducer. The imaging
feasibility of the PAE probe based on lens-focused ring shape transducer has been
demonstrated through in vivo volumetric imaging of rabbit esophageal tract. The self-
focused ring shape transducer, capable of focusing laser light beam by inner installed
92

optical lens, can be used to obtain optical-resolution PA imaging. Small aperture self-
focused circular shape transducer was employed in a shaft-based PAE probe to improved
flexibility and imaging capabilities by being inserted in the instrument channel of a
standard video endoscope.
Furthermore, a 40-MHz 32-element kerfless phased array using PMN-PT single
crystal and a 30-MHz 32-element kerfed phased array using PZT-5H ceramic were
modeled, fabricated and tested. Kerfless array’s imaging capability has been demonstrated
by US wire phantom imaging and PA graphite rod imaging. As expected, the kerfless array
directivity suffered from high crosstalk between elements. Based upon the results presented,
the kerfless array technology should be able to provide an alternate for high-frequency
phased array designs. The miniature active area of both arrays makes them as potential
candidates for PAE application.
7.2 Suggestions for Future Work
The array work presented in this dissertation is for a demonstration of principle and
realization of concepts. Further development is required to enhance the performance of
current transducers.
As discussed in Chapter 6, although the kerfless PMN-PT array exhibits very good
sensitivity, it suffers high cross talks between adjacent elements. However, the very small
dimensions of pillars/kerfs required for high frequency arrays cannot be accomplished on
PMN-PT single crystal by the traditional dicing-and-filling method due to its brittle nature.
Moreover, the kerfed PZT array shows low sensitivity because of large electrical
93

impedance mismatch. Therefore, new materials and new techniques are highly desired for
improving performance of PAE phased array.
Piezoelectric composite materials have many attractive characteristics for
ultrasonic transducer applications, especially for array transducers. For 1-3 composite
material, piezoelectric pillars was embedded in soft epoxy matrices, thus minimizing lateral
clamping between active piezoelectric elements. As a result, electromechanical coupling
coefficient of the composite is higher than that of the bulk material. Furthermore, this
special composite structure can significantly decrease inter-element crosstalk, which is
very critical for high-frequency array imaging. With the epoxy filled into kerf between
piezoelectric pillars, the acoustic impedance of the composite is lower than that of bulk
material. Hence, better acoustic impedance matching between the transducer and human
tissue can be achieved which improves energy transmitting and receiving. Consequently,
bandwidth of transducer can be enhanced and imaging resolution can be improved.
Fabrication of composites for high-frequency arrays requires that the pillars/kerfs have
very small dimensions. However, it is very challenging for the traditional dicing-and-filling
method (Cannata et al. 2006) to achieve. Microfabrication and MEMS techniques offer an
alternative for the production of high-frequency composite materials and development of
high-frequency ultrasonic arrays. Therefore PMN-PT/epoxy 1-3 composites prepared by
microfabrication processes, combining photolithography, and DRIE techniques (Liu et al.
2012) can be used for miniaturize high-frequency PAE phased arrays in the future.
In the current design, the flex circuit was bonded over the array front surface. For
high-frequency arrays, electrical connection to each element is better bonded on the back
94

side of the transducer, which simplifies the top surface topology. Thus, a better electrical
connection method also needs to be explored.

95

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Abstract (if available)

Abstract Photoacoustic imaging is a novel hybrid imaging technique that combines the virtues of both optics and ultrasound by providing the high contrast of optical imaging, while retaining the high resolution and deep depth imaging capabilities of ultrasonic imaging. Because of these advantages, recently there is much interest in developing novel photoacoustic imaging modalities. ❧ Ultrasonic transducer is very critical among all components of photoacoustic imaging system and can determine photoacoustic image quality. Similar to ultrasound imaging, the resolution and imaging depth of photoacoustic imaging is scalable, depending on characteristics of the ultrasonic transducer used, such as frequency, bandwidth and shape. ❧ This research investigated ultrasonic transducers for two photoacoustic imaging modalities: photoacoustic endoscopy and photoacoustic microscopy. Different types of high-frequency single element and array ultrasonic transducers have been designed and fabricated for each modality. Based on the specially designed ultrasonic transducers, photoacoustic endoscopy imaging probe and photoacoustic microscopy imaging system have been built. Ex vivo and in vivo studies have been carried out to validate these new imaging concepts. The imaging results strongly supported the enhancements brought by the new transducers and also suggested potential improvements in future.

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Asset Metadata

Creator Chen, Ruimin (author)

Core Title High-frequency ultrasonic transducers for photoacoustic applications

School Viterbi School of Engineering

Degree Doctor of Philosophy

Degree Program Biomedical Engineering

Publication Date 02/18/2014

Defense Date 12/10/2013

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

Tag high-frequency,OAI-PMH Harvest,photoacoustic imaging,piezoeletric material,single crystal,ultrasonic tranducer

Format application/pdf (imt)

Language English

Contributor Electronically uploaded by the author (provenance)

Advisor Shung, Kirk Koping (committee chair), Chen, Yong (committee member), Yen, Jesse T. (committee member), Zhou, Qifa (committee member)

Creator Email chen14404@hotmail.com,ruiminch@usc.edu

Permanent Link (DOI) https://doi.org/10.25549/usctheses-c3-363382

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