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Transducers and signal processing techniques for simultaneous ultrasonic imaging and therapy
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Transducers and signal processing techniques for simultaneous ultrasonic imaging and therapy
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Content
TRANSDUCERS AND SIGNAL PROCESSING TECHNIQUES FOR
SIMULTANEOUS ULTRASONIC IMAGING AND THERAPY
by
Jong Seob Jeong
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 2010
Copyright 2010 Jong Seob Jeong
ii
DEDICATION
With genuine thanks to my beloved parents, Sang Jun Jeong and Kyung Won Lee;
my precious sister, Sun Hee Jeong
for their endless love and unflagging support
iii
ACKNOWLEDGEMENTS
During five years of doctoral study in NIH Resource Center for Medical
Ultrasonic Transducer Technology at University of Southern California, I could hone my
research and development skills with the help of numerous outstanding individuals.
In the first place, I would like to give my heartily thanks to my adored parents
and precious sister for their endless love and unconditional devotion. I would like to
show my deepest appreciation to my supervisor Dr. K. Kirk Shung whose warm
encouragement, excellent support, and thoughtful guidance from the initial to end of this
thesis. Dr. Shung provided me a precious opportunity to inspire and enrich my growth as
a Ph. D. I would like to express my deeply-felt thanks to Dr. Jonathan M. Cannata for his
instructive advices and timely suggestions. Dr. Cannata has significantly enhanced my
research ability and challenged my thinking.
I gratefully acknowledge my committee members Dr. Manbir Singh, Dr. Ellis
Meng, and Dr. Eun Sok Kim for their insightful comments to allow me to finish this
thesis. I am indebted to many of my colleagues to support me. I would like to give a
special thanks to Jay A. Williams for lots of discussion and help about fabrication of
transducers. I would like to thank Dr. Jung Woo Lee, Dr. Hyung Ham Kim, and Dr. Jin
Ho Chang for their encouragement and warm advice in my graduate study and life. Many
thanks to Dr. Ruibin Liu and Dr. Gin-Shin Chen who offered me a great chance to study
HIFU transducers. I would also like to thank Dr. Chi Hyung Seo, Dr. Samer Awad, Mr.
iv
Hamid R. Chabok, Mr. Jin Hyoung Park, and Mr. Chang Yang Lee, for their kind advice
and help in my graduate experience. I thank you all from the bottom of my heart.
v
TABLE OF CONTENTS
DEDICATION ........................................................................................................................ ii
ACKNOWLEDGEMENTS ........................................................................................................ iii
LIST OF TABLES ................................................................................................................. vii
LIST OF FIGURES ............................................................................................................... viii
ABSTRACT ....................................................................................................................... xvii
CHAPTER 1: THERAPEUTIC ULTRASOUND ..........................................................................1
1.1 Introduction ............................................................................................1
1.2 HIFU with Image Guidance ...................................................................4
1.3 HIFU Treatment Time ...........................................................................9
1.4 The Scope of Research.........................................................................13
CHAPTER 2: CODED EXCITATION WITH FIXED NOTCH FILTERING FOR SIMULTANEOUS
THERAPY AND IMAGING ...............................................................................14
2.1 Introduction ..........................................................................................14
2.2 Transducer Design ...............................................................................17
2.3 Signal Processing Techniques..............................................................22
2.3.1 Acoustic Intensity ....................................................................22
2.3.2 Coded Excitation Technique ...................................................23
2.3.3 Fixed Notch Filtering Technique ............................................36
2.4 Simulations ..........................................................................................41
2.4.1 Point Target Simulation ...........................................................41
2.4.2 Intensity Simulation ................................................................48
2.5 Prototype Transducer ...........................................................................53
2.5.1 Modeling .................................................................................53
2.5.2 Fabrication ...............................................................................58
2.5.3 Performance Evaluation ..........................................................59
2.6 Experiments .........................................................................................62
2.6.1 Temperature Profile .................................................................62
2.6.2 Single Scanline Experiment ....................................................65
2.6.3 In vitro B-mode Experiment ....................................................70
2.7 Discussion ............................................................................................73
vi
CHAPTER 3: ADAPTIVE SUPPRESSION OF HIFU INTERFERENCE FOR SIMULTANEOUS
THERAPY AND IMAGING ...............................................................................77
3.1 Introduction ..........................................................................................77
3.2 Signal Processing Techniques..............................................................79
3.2.1 Fixed Notch Filtering Technique ............................................79
3.2.2 Adaptive Noise Canceling Technique .....................................84
3.3 Transducer Design and Fabrication .....................................................87
3.4 Experiments .........................................................................................90
3.4.1 Temperature Profile .................................................................90
3.4.2 Single Scanline Experiment ....................................................92
3.4.3 In vitro B-mode Experiment ....................................................94
3.5 Discussion ............................................................................................98
CHAPTER 4: DUAL-FOCUS HIFU TRANSDUCER FOR EXTENDED TISSUE LESIONS ..........101
4.1 Introduction ........................................................................................101
4.2 Methods..............................................................................................104
4.2.1 Transducer Design .................................................................104
4.2.2 Sound Field Simulation .........................................................107
4.2.3 Transducer Fabrication ..........................................................115
4.3 Experiments .......................................................................................117
4.3.1 Electrical Impedance Measurement ......................................117
4.3.2 Transmit Response Measurement .........................................118
4.3.3 DOF/Lateral Beamwidth/Sidelobe Measurement .................121
4.3.4 Bio-Heat Transfer Simulation and In vitro Experiment ........124
4.4 Discussion ..........................................................................................129
CHAPTER 5: SUMMARY AND FURTHER WORKS ..............................................................133
5.1 Summary ............................................................................................133
5.2 Further Works ....................................................................................137
BIBLIOGRAPHY .................................................................................................................139
vii
LIST OF TABLES
Table 2.1: Properties of different piezoelectric materials. ............................................ 20
Table 2.2: Properties for the therapeutic and imaging array. ........................................ 42
Table 2.3: Modeling parameters for the prototype transducer. ..................................... 55
Table 2.4: Property comparison of the 1-3 composite piezoelectric materials. ............ 56
Table 4.1: Specification of the DFTUT for sound field simulation. ........................... 109
Table 4.2: Simulated -6 dB DOF, -6 dB lateral beamwidth, and sidelobe of the
transducers. ................................................................................................ 115
Table 4.3: Parameters for fabrication of the prototype DFTUT. ................................ 116
Table 4.4: Parameters for bio-heat transfer simulation. .............................................. 125
viii
LIST OF FIGURES
Figure 1.1: Comparison of the acoustic stacks of the imaging (a) and that of
the therapeutic transducer (b)...................................................................... 2
Figure 1.2: Generation of the hyperechoic region for the ablated lesion in the
B-mode image during HIFU treatment (Owen et al., 2006). ...................... 6
Figure 1.3: Generation of the hypoechoic region for the ablated lesion (a)
before and (b) after HIFU treatment (Chan et al., 2002). ........................... 6
Figure 1.4: High amplitude of the HIFU interference signals received by the
imaging transducer (Vaezy et al., 2001b). .................................................. 8
Figure 1.5: Conventional integrated HIFU/imaging transducer based on large
frequency difference between two transducers (courtesy by EDAP
Technomed Inc.). ........................................................................................ 8
Figure 1.6: Generation of the hyperechoic region for the ablated lesion: (a)
Before and (b) during HIFU treatment (Vaezy et al., 2001b). .................... 9
Figure 1.7: Multi-step HIFU treatment process for the malignant prostate
tissue (Chinn, 2005). ................................................................................. 10
Figure 1.8: Generation of multi-foci using the array transducer (Ebbini and
Cain, 1989). ............................................................................................... 11
Figure 1.9: Generation of the broad lesion using the toric transducer
(Melodelima et al., 2009). ......................................................................... 12
Figure 1.10: Comparison of (a) single focusing and (b) split-beam focusing
schemes (Seip et al., 2001). ...................................................................... 12
ix
Figure 2.1: Simplified schematic diagram of the IMCPA transducer with 5 ×
3 elements: (a) The front view of the IMCPA without a matching
layer of imaging array. (b) The side view of the IMCPA with a
matching layer of imaging array. All arrays have 1-3
piezocomposite structures and their surfaces have a common focal
point in elevational direction. The center imaging array has a
backing layer to increase the bandwidth and the outer therapy
arrays have air backings to maximize transmission of ultrasound. A
matching layer would also increase the transmission efficiency of
imaging array. ........................................................................................... 19
Figure 2.2: (a) Received 2-cycle short pulses for imaging without CW
interference, (b) with CW interference, (c) frequency response of
(b), and (d) envelope signal of (b). ........................................................... 23
Figure 2.3: Signal processing for real-time imaging during therapy by using
the IMCPA. (a) The therapeutic and coded imaging signals are
emitted to the target at the same time. (b) The reflected therapeutic
signal received by the imaging array can be removed by means of
notch filtering and pulse compression. ..................................................... 25
Figure 2.4: (a) 13-bit Barker code with 1 cycle per bit, (b) frequency response
of (a), (c) after pulse compression, and (d) after conventional
sidelobe suppression filtering. .................................................................. 27
Figure 2.5: (a) 13-bit Barker code with 2 cycles per bit, (b) frequency response
of (a), (c) after pulse compression, and (d) after conventional
sidelobe suppression filtering. .................................................................. 28
Figure 2.6: (a) 13-bit Barker code with 3 cycles per bit, (b) frequency response
of (a), (c) after pulse compression, and (d) after conventional
sidelobe suppression filtering. .................................................................. 29
Figure 2.7: Transducer impulse responses modeled by a 4
th
order Butterworth
filter for point target imaging simulation: (a) Imaging transducer
with 50 % -6 dB bandwidth and (b) therapeutic transducer with
30 % -6 dB bandwidth. ............................................................................. 30
x
Figure 2.8: (a) Received 13-bit Barker code with 2 cycles per bit without CW
interference, (b) with CW interference, (c) frequency response of
(b), and (d) envelope signal of (b) after pulse compression with
sidelobe suppression filter......................................................................... 31
Figure 2.9: (a) Received 13-bit Barker code with 3 cycles per bit without CW
interference, (b) with CW interference, (c) frequency response of
(b), and (d) envelope signal of (b) after pulse compression with
sidelobe suppression filter......................................................................... 32
Figure 2.10: (a) Time domain response of the Dolph-Chebyshev-windowed
chirp signal, (b) frequency response of (a), and (c) envelope signal
of (a) after pulse compression. .................................................................. 34
Figure 2.11: (a) Received Dolph-Chebyshev-windowed chirp signal without
CW interference, (b) with CW interference, (c) frequency response
of (b), and (d) envelope signal of (b) after pulse compression with
a mismatched filter. ................................................................................... 35
Figure 2.12: Frequency responses of 4 MHz and 8 MHz notch filters. The notch
attenuation values are -37 dB and -31 dB at 4 MHz and 8 MHz,
respectively. .............................................................................................. 38
Figure 2.13: Envelopes of imaging signals with CW interference signals: 2-
cycle short pulses, (a) before and (b) after notch filtering; the 13-
bit Barker code with 2 cycles per bit, (c) before and (d) after notch
filtering; the 13-bit Barker code with 3 cycles per bit, (e) before
and (f) after notch filtering. ....................................................................... 39
Figure 2.14: Envelopes of the Dolph-Chebyshev-windowed chirp signals with
CW interference signals: Before (dashed line) and after (solid line)
notch filtering. ........................................................................................... 40
Figure 2.15: Geometrical surface of the therapeutic array aperture combined
with a mask window of the aperture. Outer two arrays are activated
for transmit mode (left) and inner array is activated for receive
mode (right). ............................................................................................. 43
xi
Figure 2.16: Geometrical surface of the imaging array combined with a mask
window of the aperture. Only inner array is activated during both
transmit and receive mode. ....................................................................... 44
Figure 2.17: Point target simulation with imaging signals in which CW
interference signals were superimposed. All figures were
logarithmically compressed with a dynamic range of 40 dB. 2-
cycle short pulses, (a) without and (b) with interference; the 13-bit
Barker code with 2 cycles per bit, (c) without and (d) with
interference; the 13-bit Barker code with 3 cycles per bit, (e)
without and (f) with interferences. ............................................................ 45
Figure 2.18: Simulated point target images of interference-mixed imaging
signals after notch filtering. All figures were logarithmically
compressed with a dynamic range of 40 dB. (a) 2-cycle short
pulses, (b) the 13-bit Barker code with 2 cycles per bit, and (c) the
13-bit Barker code with 3 cycles per bit. .................................................. 47
Figure 2.19: Relative intensity plot of the therapeutic array: (a) Contour plot,
(b) azimuthal profile, and (c) axial profile. ............................................... 49
Figure 2.20: Relative intensity plot of the imaging array: (a) Contour plot, (b)
azimuthal profile, and (c) axial profile. .................................................... 50
Figure 2.21: Pressure plot of the therapeutic array: (a) Axial profile and (b)
contour plot. .............................................................................................. 51
Figure 2.22: Intensity (I
SPTA
) plot for the therapeutic array: (a) Axial profile and
(b) contour plot. ........................................................................................ 52
Figure 2.23: 3D plot of the intensity (I
SPTA
) for the therapeutic array. .......................... 53
Figure 2.24: Dimension of (left) therapeutic and (right) imaging transducer. All
transducers are composed of 1-3 piezocomposite. ................................... 57
Figure 2.25: Acoustic stack of (left) therapeutic and (right) imaging transducer. ........ 57
xii
Figure 2.26: (a) Photograph of the prototype integrated HIFU/imaging
transducer and (b) cross sectional schematic diagram. ............................. 59
Figure 2.27: Electrical impedance of therapeutic transducer using KLM
simulation and an impedance analyzer: (a) Magnitude and (b)
phase. ........................................................................................................ 60
Figure 2.28: Electrical impedance of imaging transducer using KLM simulation
and an impedance analyzer: (a) Magnitude and (b) phase. ....................... 60
Figure 2.29: Pulse echo response of the therapeutic transducer in the frequency
domain....................................................................................................... 61
Figure 2.30: Pulse echo response of the imaging transducer in the frequency
domain....................................................................................................... 62
Figure 2.31: (a) Photograph and (b) schematic diagram for temperature
measurement with the prototype integrated HIFU/imaging
transducer. ................................................................................................. 63
Figure 2.32: (a) Measured temperature profile and (b) coagulated sliced porcine
muscle using the prototype integrated HIFU/imaging transducer
with spherical focusing. ............................................................................ 64
Figure 2.33: (a) Photograph and (b) schematic diagram for a single scanline
experiment to test the performance of the prototype integrated
HIFU/imaging transducer. ........................................................................ 66
Figure 2.34: Measured envelope signals by using the prototype transducer: (a)
2-cycle short pulse excitation before notch filtering, (b) 2-cycle
short pulse excitation after notch filtering, (c) the 13-bit Barker
code excitation with 2 cycles per bit excitation after notch filtering
and pulse compression, (d) the 13-bit Barker code excitation with 3
cycles per bit after notch filtering and pulse compression. The
second peak in the displayed data at approximately 21 μs is the
reflected signal from the bottom of the target. .......................................... 68
xiii
Figure 2.35: Envelope signals of the Dolph-Chebyshev-windowed chirp signal
(dashed line) before and (solid line) after notch filtering. ........................ 70
Figure 2.36: Photograph with a schematic diagram for single element imaging
with the activated HIFU transducer. ......................................................... 71
Figure 2.37: B-mode images with a slice of porcine muscle after pulse
compression: (a) Original image of the Dolph-Chebyshev-
windowed chirp. After mixing HIFU interference (b) before and
(c) after notch filtering. ............................................................................. 72
Figure 3.1: Frequency response of the fixed notch filters designed by 2
nd
order
IIR type at 4 MHz and 8 MHz frequency. The double arrow
indicates the -3 dB bandwidth of the notch filter. ..................................... 82
Figure 3.2: Simulated envelope signals for the fixed notch filters: (a) Original
2-cycle short-pulse signal. (b) When the therapeutic interference
signals are mixed with (a). (c) After using the fixed notch filter
with Q=8/16. (d) After using the fixed notch filter with Q=2/4. .............. 83
Figure 3.3: Block diagram for (a) the simultaneous therapy and imaging
system and (b) adaptive noise canceling algorithm. ................................. 86
Figure 3.4: Simulated envelope signals for the adaptive noise canceling
technique (solid line) and the original signal (dashed line). ..................... 87
Figure 3.5: Schematic diagram for the prototype integrated HIFU/imaging
transducer with cylindrically focusing: (a) A photograph and (b) a
cross sectional drawing of the side view................................................... 88
Figure 3.6: (a) Schematic diagram for temperature measurement for the
prototype integrated HIFU/imaging transducer. (b) Measured
temperature profile in 60 seconds on a slice of porcine muscle. .............. 91
xiv
Figure 3.7: (a) Photograph and (b) schematic diagram for a plate experiment
for testing the performance of the prototype integrated
HIFU/imaging transducer. ........................................................................ 92
Figure 3.8: Measured pulse echo envelope signals obtained by the prototype
integrated HIFU/imaging transducer: (a) Original 2-cycle short-
pulse excitation. (b) When the reflected therapeutic interference
was mixed with (a) received by imaging transducer. (c) After notch
filtering with a Q being 8 and 16 as shown in Figure 3.1. (d) After
adaptive noise canceling. .......................................................................... 94
Figure 3.9: (a) Photograph and (b) schematic diagram for imaging during
HIFU emission with the prototype integrated HIFU/imaging
transducer. ................................................................................................. 95
Figure 3.10: (a) Original image of a slice of porcine muscle, (b) after activating
HIFU transducer, (c) after notch filtering, and (d) after adaptive
noise canceling. All figures were logarithmically compressed with
a dynamic range of 40 dB. ........................................................................ 97
Figure 4.1: Schematic diagram of the dual-focus therapeutic ultrasound
transducer (DFTUT) with specification: (a) Side view and (b) front
view. Note that the relative geometric focus offset between two
focal points is 5.24 mm. .......................................................................... 105
Figure 4.2: Schematic diagram of the DFTUT aperture used in the Field-II
simulation. The radii of curvatures for inner and outer elements are
19 mm and 24 mm, respectively. The diameters of inner and outer
elements are 12 mm and 21 mm, respectively. ....................................... 108
Figure 4.3: Transmit beam profile of the inner element with 19 mm focal
depth: (a) A contour plot in the decibel scale, (b) lateral beam
profile, and (c) axial beam profile. .......................................................... 111
Figure 4.4: Transmit beam profile of the outer element with 24 mm focal
depth: (a) A contour plot in the decibel scale, (b) lateral beam
profile, and (c) axial beam profile. .......................................................... 112
xv
Figure 4.5: Transmit beam profile of the DFTUT: (a) A contour plot in the
decibel and (b) axial beam profile. ......................................................... 113
Figure 4.6: Transmit beam profile of the single focused transducer with 21.5
mm focal depth: (a) A contour plot in decibel, (b) lateral beam
profile, and (c) axial beam profile. .......................................................... 114
Figure 4.7: Photograph of the prototype DFTUT. ..................................................... 117
Figure 4.8: Measured electrical impedance of the DFTUT with a water load.
There are two impedance peaks in series for inner and outer
elements. ................................................................................................. 118
Figure 4.9: Experimental setup for measurement of the transmit response,
DOF, and lateral beamwidth of the DFTUT by using a hydrophone:
(a) A photograph and (b) a schematic diagram. ...................................... 119
Figure 4.10: Frequency domain plots of the measured transmit response along
the axial direction: (a) 18 mm, (b) 23 mm, (c) 28 mm, and (d) 33
mm in depth. ........................................................................................... 120
Figure 4.11: Simulated and measured data for the DFTUT using a hydrophone:
(a) An axial beam profile with DOF and (b) -6 dB overall lateral
beamwidth within the -6 dB DOF. .......................................................... 123
Figure 4.12: Simulated and measured lateral beam pattern for the DFTUT: The
simulated (red-solid line) and the measured data at 20 mm (black-
solid line) and at 22 mm in depth (blue-dashed line). Note that the
pedestal level of the all measured data is higher than simulated
data due to ADC noise. ........................................................................... 124
Figure 4.13: Simulated temperature distribution for the DFTUT. The position of
the transducer was on the left side as indicated by a depth of zero. ....... 128
xvi
Figure 4.14: Cross-section of a piece of beef liver after HIFU sonication with
the DFTUT. The arrow indicates that the HIFU exposure direction.
(a) A schematic diagram to generate (b) the ablated lesion in a beef
liver specimen by using the DFTUT. ...................................................... 129
Figure 5.1: Schematic diagram for the integrated multi-functional confocal
phased array (IMCPA). ........................................................................... 138
xvii
ABSTRACT
Recently, high intensity focused ultrasound (HIFU) was successfully used for
noninvasive treatment of the benign or malignant tissues. In ultrasound image-guided
HIFU (US-gHIFU) using an integrated HIFU/imaging transducer, real-time simultaneous
therapy and imaging is more desirable because it allows for tracking tissue movement
and monitoring feedback induced by a treated target. However, reflected HIFU signals
corrupt the quality of signals received by an imaging transducer during treatment.
In the first part of this thesis, it was demonstrated that these interference signals
can be significantly reduced in the formed brightness mode (B-mode) ultrasound image
by implementing coded excitation with fixed notch filtering. In the second part, short
pulse excitation with adaptive noise canceling technique was proposed to optimally
suppress therapeutic interference with variable amplitudes while maintaining the original
image form as closely as possible. To demonstrate the performance of these techniques, a
design of the integrated HIFU/imaging phased array transducer was proposed for
treatment of malignant prostate tissues. The center row forms imaging signals and the
two identical outer rows work together to produce HIFU signals. As a preliminary
experiment, the prototype integrated HIFU/imaging transducer composed of three single
elements was built and the performance of the proposed techniques was verified with a
soft biological tissue specimen.
xviii
The third part investigates formation of a large tissue lesion per HIFU sonication
to reduce the overall treatment time. The goal of this study is to show the feasibility of
enlarging tissue lesion size with a dual-focus therapeutic ultrasound transducer by
increasing the depth-of-focus (DOF). The proposed transducer for treatment of the
malignant prostate tissues consists of a disc- and an annular-type element of different
radii of curvatures to produce two focal zones. To compensate attenuation and to
maintain uniform beamwidth of the elongated DOF, each element transmits ultrasound of
a different center frequency: the inner element at a higher frequency for near field
focusing and the outer element at a lower frequency for far field focusing. By activating
two elements at the same time with a single transmitter capable of generating a dual-
frequency mixed signal, the overall ablated region of the proposed transducer can be
considerably extended. By employing this design to integrated HIFU/imaging transducer,
not only real-time simultaneous therapy and imaging, but also reduced HIFU treatment
time can be achieved at the same time.
1
CHAPTER 1: THERAPEUTIC ULTRASOUND
1.1 Introduction
Ultrasound whose frequency is higher than 20 KHz has been widely used for
non-destructive testing, medical ultrasonic imaging, and ultrasound therapy. Typically,
ultrasound in the 2 MHz − 15 MHz frequency range is used for diagnostic ultrasound and
especially 1 MHz − 4 MHz frequency range is preferred for ultrasound treatment of the
malignant tissue considering target distance, intensity, attenuation, and ablated lesion size.
In general, an ultrasound imaging transducer consists of a matching layer, a
piezoelectric layer, and a backing layer. A piezoelectric layer converts mechanical energy
to electrical energy and vice versa. A matching layer can increase the sensitivity and the
bandwidth by reducing acoustic impedance mismatch between a piezoelectric material
and a medium. A backing layer suppresses ringing of the ultrasound and thus results in a
broad bandwidth. In the case of therapeutic ultrasound transducer, its configuration
should be different from that of the imaging transducer. The heating on the surface of the
piezoelectric material may cause detachment of the matching layer from the transducer,
and a backing layer may sacrifice sensitivity of the ultrasound during high voltage
operation. Therefore, it is desirable to fabricate a therapeutic transducer by merely using
a piezoelectric material as shown in Figure 1.1.
2
(a) (b)
Figure 1.1: Comparison of the acoustic stacks of the imaging (a) and that of the
therapeutic transducer (b).
When a short pulse signal is applied to the imaging transducer, the intensity at
the focal point is low enough to protect the normal tissue while therapeutic transducer
driven by a high amplitude sinusoidal wave can destroy tissues at the focal point. It is
well known that two mechanisms affect generation of the tissue lesion. One is thermal
effect induced by high temperature and the other is mechanical effect due to cavitation
during treatment. Although cavitation may disturb propagation of the ultrasound, it can
enhance ablation effect based on mechanical stress of bubbles, which have been
demonstrated experimentally (Kennedy et al., 2003).
There are two kinds of ultrasound therapeutic methods depending on the
operating temperature. One is hyperthermia by maintaining temperature of the target
from 43 °C to 45 °C in 60 minutes and thus arrests reproduction of tissue (Kennedy et al.,
2003). The other is high intensity focused ultrasound (HIFU) by increasing tissue
temperature more than 70 °C in a short time to make coagulated necrosis, i.e., irreversible
3
cell death (Armour et al., 1993). To obtain HIFU treatment effect, more than 1000 W/cm
2
of the spatial-peak temporal-average intensity (I
SPTA
) should be obtained at the focal area.
Note that I
SPTA
of diagnostic ultrasound should be less than 100 mW/cm
2
based on the
food and drug administration (FDA) guidance (Barnett et al., 2000; ter Haar et al., 1995).
It has been also demonstrated that the continuous wave (CW) may provide better
treatment performance for HIFU comparing to pulsed wave (PW) with a low duty factor
(Daum and Hynynen, 1999b; Hundt et al., 2008).
Typically, the performance of the HIFU transducer is measured by a transducer
efficiency as shown in the below equation.
electric acoustic
P P η = (1.1)
where P
acoustic
is the total acoustic power (TAP) and P
electric
is the loading electrical power,
η is the transducer efficiency. TAP can be measured by a radiation force balance
equipment or by using a hydrophone based on the below formula.
2
8 . 1 867 . 0
1
D
I
P
SP
acoustic
= (1.2)
where I
SP
is the spatial peak intensity at focal point and D is the beamwidth of the -6 dB
of the maximal pressure (Chen et al., 1998; Hill et al., 1994; Köhrmann et al., 2002;
Malcolm et al., 1996).
4
On the other hand, I
SPTA
can be calculated using measured pressure, impedance
of the medium, and the pulse repetition time (PRT).
df
m
SPTA
t
c
P
I
=
ρ 2
2
(1.3)
where P
m
is the pressure measured at focal point, ρ is the density, c is the velocity, and t
df
is the duty factor.
1.2 HIFU with Image Guidance
In recent years, HIFU has become increasingly important in the noninvasive
treatment of malignant tissues due to a relatively short treatment time, a small destroyed
tissue area, and a fast recovery time compared to open surgery (Vaezy, 2001a). Several
clinical studies have been conducted to investigate the feasibility of HIFU treatment for
breast (Furusawa et al., 2007), liver (Daum et al., 1999a), and prostate cancer (Azzouz et
al., 2006; Blana et al., 2004; Sanghvi et al., 1999).
In noninvasive HIFU surgery, the treatment region is first scanned with a medical
imaging modality such as magnetic resonance imaging (MRI), computer tomography
(CT), or ultrasound to assist in mapping tissue pathologies prior to treatment. Typically,
precise targeting without damaging healthy tissues and real-time monitoring the response
of the treated target is of critical importance to a physician for efficient HIFU surgery
5
(Vaezy et al., 2001a; Wu et al., 2008). To satisfy these requirements, the imaging
modality should provide real-time visualization simultaneously during treatment. MRI
provides a high-resolution image and an efficacious temperature map, but it is expensive
and requires a large space. Ultrasound can offer advantages in real-time imaging, cost-
effectiveness, excellent portability, and potential integration with other devices (Ebbini et
al., 2006). Currently, commercialized ultrasound image guided HIFU (US-gHIFU) has
been widely used for treatment of the localized prostate cancer (Azzouz et al., 2006;
Blana et al., 2004; Chinn, 2005).
There are several ultrasound imaging approaches that have been used in US-
gHIFU such as the B-mode imaging, elastography, and radiation force imaging (Chan et
al., 2002; Konofagou and Hynynen, 2003; Melodelima et al., 2007; Owen et al., 2006;
Righetti et al., 1999). In B-mode imaging a hyperechoic region may be observed
generated by micro-bubbles in the targeted tissue during treatment as shown in Figure 1.2.
After treatment, a hypoechoic region may be observed resulted from different attenuation
or reflection coefficients related to the hardness variation between the coagulated target
and normal tissues (Chan et al., 2002; Owen et al., 2006; Vaezy et al., 2001b).
6
Figure 1.2: Generation of the hyperechoic region for the ablated lesion in the B-mode
image during HIFU treatment (Owen et al., 2006).
(a) (b)
Figure 1.3: Generation of the hypoechoic region for the ablated lesion (a) before and (b)
after HIFU treatment (Chan et al., 2002).
7
One of the popular methods to realize real-time simultaneous ultrasound therapy
and imaging is to combine a HIFU transducer with an ultrasound imaging transducer
(Fleury et al., 2006). However, when the HIFU transducer is activated, strong HIFU
noises corrupt ultrasound image quality generated by imaging transducer as shown in
Figure 1.4. Thus, several techniques have been used in attempting to reduce the strong
reflected HIFU interference signals received by an imaging transducer. Notably, using a
large frequency difference between HIFU and imaging transducers in Figure 1.5 (Azzouz
et al., 2006, Kluiwstra et al., 1997), or interleaving the HIFU data with the imaging data
as shown in Figure 1.6 (Bouchoux et al., 2008; Seip et al., 2002; Vazey et al., 2001b)
have all shown promise in the reduction of the influence of HIFU signal on the image.
Figure 1.5 shows the conventional technique using large frequency difference between
HIFU and imaging transducer to avoid HIFU interference harmonics. Although this type
transducer can reduce HIFU interference signals, it may provide low resolution imaging
due to narrow available bandwidth about 40 %. Interleaving HIFU data with the imaging
data usually has been shown the interference region related to the frame rate as shown in
Figure 1.6(b).
8
Figure 1.4: High amplitude of the HIFU interference signals received by the imaging
transducer (Vaezy et al., 2001b).
Figure 1.5: Conventional integrated HIFU/imaging transducer based on large frequency
difference between two transducers (courtesy by EDAP Technomed Inc.).
9
Figure 1.6: Generation of the hyperechoic region for the ablated lesion: (a) Before and
(b) during HIFU treatment (Vaezy et al., 2001b).
1.3 HIFU Treatment Time
There is another drawback in current HIFU transducer. The total HIFU treatment
time is relatively long, i.e., more than from 3 to 4 hours depending on the target size. The
most dominant factor causing this problem is a narrow region of -6 dB intensity (ter Haar,
2000). Typically, the size of the ablated lesion related to -6 dB of the maximal intensity is
10
too small and thus multi-step HIFU treatment has been used for treatment of the large
target as shown in Figure 1.7 (Chinn, 2005).
Figure 1.7: Multi-step HIFU treatment process for the malignant prostate tissue (Chinn,
2005).
To figure out this problem, several techniques were developed. One of them is
using an array type transducer based on multi-subaperture activation to generate multi-
foci as shown in Figure 1.8. But it is required the complicated transducer fabrication and
system.
11
Figure 1.8: Generation of multi-foci using the array transducer (Ebbini and Cain, 1989).
Another method is using the toric transducer to generate a broad lesion in the
lateral and elevational direction as shown in Figure 1.9. Also, the split-beam focusing
technique in Figure 1.10 can extend the ablated lesion by using the subdivided aperture
with sectional electrodes. Note that most techniques provide the broad lesion in the lateral
and elevational direction.
12
Figure 1.9: Generation of the broad lesion using the toric transducer (Melodelima et al.,
2009).
(a) (b)
Figure 1.10: Comparison of (a) single focusing and (b) split-beam focusing schemes
(Seip et al., 2001).
13
1.4 The Scope of Research
To provide real-time imaging during treatment especially targeting malignant
prostate tissues, novel integrated HIFU/imaging transducers and signal processing
techniques were proposed in this thesis. This research shows the proposed integrated
HIFU/imaging transducers can provide maximal therapy and imaging performance based
on different material and configuration. Subsequently, it was demonstrated that the HIFU
interference received by the imaging transducer can be minimized using not only coded
excitation with fixed notch filtering, but also short pulse excitation with adaptive noise
cancellation. It was shown that the dual-focus HIFU transducer can decrease treatment
time by increasing ablated lesions in the axial direction, and thus it can be applied for the
integrated HIFU/imaging transducer.
This thesis consists of five chapters. Chapter 1 describes general background of
HIFU about used frequency, transducer configuration, bio-effect, application, intensity,
image guidance, and treatment time. Chapter 2 and Chapter 3 discuss coded excitation
with fixed notch filtering and short pulse excitation with adaptive noise cancellation
method, respectively. In Chapter 4, dual-focus transducer was illustrated. Chapter 5
summarizes all results in this research with potential further works.
14
CHAPTER 2: CODED EXCITATION WITH FIXED NOTCH FILTERING
FOR SIMULTANEOUS THERAPY AND IMAGING
2.1 Introduction
Recently, ultrasound image-guided HIFU (US-gHIFU) has been widely used for
treatment of a localized prostate cancer and shown its feasibility (Azzouz et al., 2006;
Blana et al., 2004). There are several ultrasound imaging modalities in US-gHIFU such
as B-mode image and radiation force imaging (Chan et al., 2002; Konofagou and
Hynynen, 2003; Melodelima et al., 2007; Owen et al., 2006). Among them, B-mode
image can provide hyperechoic region due to bubble formation during treatment. After
treatment, it can display hypoechoic region due to attenuation changes, or its stiffness
variations (Chan et al., 2002; Melodelima et al., 2007; Owen et al., 2006; Vaezy et al.,
2001b). In noninvasive HIFU surgery, precise targeting without damage of normal tissues
and real-time monitoring the response of the treated target is important to a physician for
efficient HIFU surgery (Wu et al., 2008). One of the popular methods to realize
simultaneous ultrasound therapy and imaging is to combine HIFU transducer with
ultrasound imaging transducer (Fleury et al., 2006). However, when the HIFU transducer
is activated, strong HIFU noise interferes ultrasound imaging generated by imaging
transducer. Thus, several techniques were used in attempt to reduce the strong reflected
therapeutic signals received by an imaging transducer. Particularly, using a large
15
frequency difference between HIFU and imaging transducers (Azzouz et al., 2006;
Kluiwstra et al., 1997), or synchronized therapeutic and imaging data (Seip et al., 2002)
have all shown some success in reduction of the HIFU signal’s influence on the image.
In a similar effort to realize simultaneous therapy and imaging, this research
describes the feasibility of real-time imaging during sonication of the high amplitude
continuous wave (CW). For this purpose, a design of the HIFU transducer called the
integrated multi-functional confocal phased array (IMCPA) was proposed. The
transducer consists of triple-row phased arrays: a 6 MHz array in the center row for
imaging and two 4 MHz arrays in the outer rows for therapy. Since one of the most
common applications by using the US-gHIFU system is currently the treatment of
prostate tissue, all design specifications such as dimension, frequency, and focal depth of
the proposed transducer targeted this application.
The most critical issue to achieve our goal with the IMCPA is how to suppress
reflected therapeutic signals received by the center-row array which is used for imaging.
When CW signals for treatment and short pulse signals for imaging are transmitted to a
target at the same time, the imaging signal may hardly be detected due to the high
amplitude of the reflected therapeutic signals. One simple way to solve this problem may
be either to decrease the intensity of transmitted therapeutic signals or to increase the
intensity of transmitted imaging signals. However, these are not practical solutions
because the intensity of therapeutic signal should be large enough to produce thermal
16
necrosis, and the intensity of diagnostic ultrasound must be below that mandated by the
food and drug administration (FDA) (Barnett et al., 2000; ter Haar et al., 1995).
As a practical solution to these limitations, this research proposes a coded
excitation technique with a notch filter to form a brightness mode (B-mode) image during
therapy. Through Field-II simulation studies (Jensen, 1996) and experimental results, it
was demonstrated that the proposed method might be used to effectively suppress the
interference signals during B-mode imaging while therapy was being carried out. The
prototype integrated HIFU/imaging transducer was built for treatment of malignant
prostate tissue. It is composed of single 6 MHz imaging element in center and two 4 MHz
HIFU elements in outer with confocal focusing in elevational direction. To maximize the
performance of the each transducer, different piezoelectric material and configuration
were used in fabrication of the transducer. The 13-bit Barker codes with 2 and 3 cycles
per bit and the Dolph-Chebyshev-windowed chirp signal were used as the coded
excitation considering axial resolution and noise robustness. Two notch filters with 4
MHz and 8 MHz notch attenuation frequency were designed using 2
nd
order infinite
impulse response filter based on a quadratic scheme (Widrow and Stearns, 1985; Winder,
2002).
Three types of experiments were conducted in this research. First, the heating
performance of the HIFU transducer was demonstrated using a thermocouple within the
in vitro porcine muscle. For algorithm test using one scanline, a strong reflector with a
polished aluminum plate was used. This experimental setup generated high amplitude of
17
HIFU interference including harmonics. In vitro experiment with a slice of porcine
muscle was conducted to obtain B-mode image. Because the imaging transducer in the
prototype integrated HIFU/imaging transducer is single element, a single element
imaging system controlled by a linear motor was used to obtain B-mode image.
2.2 Transducer Design
A HIFU array transducer should provide highly flexible control to change a focal
spot along either lateral or axial direction without moving the transducer (Ebbini et al.,
1989; Fan et al., 1995). There have been several types of multi-row array transducers for
US-gHIFU system reported in the literature (Ishida et al., 2003; Stephens et al., 2008).
The IMCPA proposed in the thesis was also a multi-row array transducer composed of
triple-row phased arrays: a 6 MHz array in the center row for imaging and two 4 MHz
arrays in the outer rows for therapy as shown in Figure. 2.1. For HIFU therapy, a
frequency range from 1 MHz to 4 MHz was preferred to increase thermal treatment effect
(Hynynen, 1991). The frequencies in the range from 3 MHz to 4 MHz have been widely
used for treatment of prostate cancer given the depth of penetration (Azzouz et al., 2006;
Blana et al., 2004; Sanghvi et al., 1999). Considering the requirements on efficient
thermal necrosis such as depth of penetration (i.e., 4 cm − 5 cm for prostate), needed
intensity more than 1000 W/cm
2
of the spatial-peak temporal-intensity (I
SPTA
) and
effective -6 dB focal area, 4 MHz and 6 MHz frequencies were respectively chosen for
18
treatment and imaging for the IMCPA. Since the second harmonic component of 4 MHz
therapeutic signal would be generated at approximately 8 MHz, the imaging transducer
should have an effective bandwidth of between 4 MHz and 8 MHz. Therefore, a 6 MHz
transducer with a -6 dB fractional bandwidth more than 50 % is required for imaging.
That is, the -6 dB bandwidth of 6 MHz array should be large enough to include the
critical frequency components of imaging signals but narrow enough to minimally
overlap with the notch frequencies.
For the purpose of efficient therapy and imaging with the IMCPA, the 6 MHz
imaging array has 64 elements with 0.6λ = 150 μm pitch, 25 μm kerf, and 8 mm height,
while each 4 MHz therapy array has 64 elements with 0.6λ = 225 μm pitch, 25 μm kerf,
and 10 mm height. The total dimension of the IMCPA is 14.4 mm × 28 mm. These
dimensions are acceptable as an endocavity transducer. The dimension of the therapy
array is 14.4 mm × 20 mm, which can generate I
SPTA
of about 1000 W/cm
2
at a focal
spot (Tavakkoli et al., 2003). The therapy array with 0.6λ spacing between elements
should be able to protect normal tissues from the grating lobes of HIFU beam. The f-
number of elevational direction is 1.4 at a focal depth of 40 mm. The novel aspect of the
IMCPA is that all arrays are focused on the same region by a geometrically curved
surface in the elevational direction eliminating the need for orienting transducers
(Stephens et al., 2008). Also, a press-focused array is more efficacious than a lens-
focused array because of acoustic energy absorption by a lens itself (Hill, 1994). Figure
19
2.1 describes a schematic diagram of the IMCPA with 5 × 3 elements as a simple
example.
Figure 2.1: Simplified schematic diagram of the IMCPA transducer with 5 × 3
elements: (a) The front view of the IMCPA without a matching layer of imaging array. (b)
The side view of the IMCPA with a matching layer of imaging array. All arrays have 1-3
piezocomposite structures and their surfaces have a common focal point in elevational
direction. The center imaging array has a backing layer to increase the bandwidth and the
outer therapy arrays have air backings to maximize transmission of ultrasound. A
matching layer would also increase the transmission efficiency of imaging array.
20
In the IMCPA design, all arrays should be fabricated with a piezoelectric 1-3
composite for surface conformation and for its low acoustic impedance. However, the
materials constituting 1-3 piezocomposite used for therapeutic transducers are typically
different from those for imaging transducers. The piezoelectric materials with a high
Curie temperature and a low dielectric/mechanical loss or high dielectric/mechanical
quality factor such as PZT4 and PZT8 are more desirable for therapeutic transducers as
described in Table 2.1 (Geng et al., 1999; Kino, 1987; Zhang et al., 2005; Zipparo et al.,
2003). PZT-5H with a high dielectric constant and electromechanical coupling has been
widely used for imaging array transducers (Fiore et al., 1996). A high thermal resistance
epoxy with a high glass transition temperature may be used for a composite therapeutic
transducer resulting in improved temperature durability. A volume fraction ratio between
piezoelectric material and epoxy of the IMCPA of 79 % was selected, given the
specifications, which should be enough to generate the required acoustic intensity (Geng
et al., 1999).
Table 2.1: Properties of different piezoelectric materials.
PZT4 PZT8 PZT-5H
a
Density, ρ [kg/m
3
] 7500 7600 7500
a
Longitudinal velocity, v
l
[m/s] 4600 4580 4560
a
Coupling coefficient, k
t
0.51 0.48 0.51
a
Relative clamped dielectric constant, ε
33
s
/ ε
0
635 580 1470
a
Mechanical quality factor, Q
m
500 1000 65
b
Curie temperature, T
c
[°C] 330 330 218
a
(Kino, 1987)
b
www.americanpiezo.com
21
A matching layer would reduce the acoustic impedance mismatch between the
transducer and the body resulting in high transmission efficiency. In the fabrication of the
HIFU transducer, the application of a matching layer may be problematic. Usually, the
area of the HIFU transducer surface is too large to maintain uniform thickness and a
uniform bonding line. Heating on the surface of the transducer may also cause
detachment of the layer from the transducer during high voltage operation. Thus, the
application of a matching layer must be carefully considered for HIFU transducer. In the
IMCPA, a matching layer was used for only imaging array. It is also advisable to
construct a HIFU transducer without a backing layer to allow maximal energy
transmission in the forward direction and to alleviate fabrication difficulties (Fan et al.,
1995; Zhang et al., 2005).
A strategy for fabricating the IMCPA is to assemble together individual arrays
after they are designed and fabricated separately. A press-focused method may be used
before or after combining these transducers. Another advantage of this configuration is
that the amplitude of reflected therapeutic signals received by the imaging transducer
might be reduced due to an elevational angle difference between the therapy and the
imaging arrays, causing most reflected therapeutic signals be directed toward the therapy
arrays rather than the imaging array.
22
2.3 Signal Processing Techniques
2.3.1 Acoustic Intensity
For therapy, I
SPTA
at a focal point should be higher than 1000 W/cm
2
to
accomplish thermal necrosis (ter Haar, 1995). The PW with a high duty cycle might be
used if it can satisfy the requirement. Although the proposed method can employs PW as
well as CW signals for therapy, a 4 MHz CW signal with a 100 % duty factor was used
so as to yield an intensity of 1000 W/cm
2
for simulations and experiments presented in
this thesis. A 6 MHz 1-cycle short pulse was chosen for imaging. Its duty factor was
0.042 % under the condition of a typical 2.5 kHz pulse repetition frequency (PRF) so that
I
SPTA
could be 18.8 mW/cm
2
as a diagnostic intensity in accordance with the FDA
guideline (Henderson et al., 1995). With these parameters, the peak pressures of the CW
signal and the 1-cycle short pulse at a focal point were computed and found to be 7.75
MPa and 1.16 MPa, respectively, from the formula (Shung, 2006):
df
SPTA
t
ZI 2
P = (2.1)
where Z is the acoustic impedance of water (1.5 MRayl) and t
df
represents the duty factor.
Given the two computed peak pressure values, transmit ultrasound pressures for both
therapy and imaging were adjusted in all simulations. It was assumed that the amplitude
of the second harmonic signal at 8 MHz was 10 dB lower than that at its fundamental
frequency (Jensen, 2002).
23
2.3.2 Coded Excitation Technique
To demonstrate proposed theory, computer simulation was conducted using
Matlab (The MathWorks Inc., Natick, MA) program. When the IMCPA fires 2-cycle
short pulses for normal imaging and CW signals for therapy, the imaging array would
receive echoes containing the high amplitude of 4 MHz and 8 MHz interference signals
as shown in Figure 2.2(b).
Figure 2.2: (a) Received 2-cycle short pulses for imaging without CW interference, (b)
with CW interference, (c) frequency response of (b), and (d) envelope signal of (b).
24
This interference decreases the signal-to-noise ratio (SNR) of the imaging signals.
The range sidelobe level of the envelope signal extracted from the echoes (Figure. 2.2(d))
was found to be approximately -4 dB, thus resulting in poor image quality. The easiest
way to improve the SNR for imaging is to increase input power of the imaging array, but
it may violate the FDA guideline. As a practical solution to this limitation, a coded
excitation technique can be used for imaging with the IMCPA because this technique can
improve the SNR by increasing the average power without changing the peak power
(O’Donnell, 1992). In addition, the correlation between coded imaging signals and CW
interference signals is significantly lower than that when a short pulse signal is used for
imaging.
25
Figure 2.3: Signal processing for real-time imaging during therapy by using the IMCPA.
(a) The therapeutic and coded imaging signals are emitted to the target at the same time.
(b) The reflected therapeutic signal received by the imaging array can be removed by
means of notch filtering and pulse compression.
Figure 2.3 depicts the proposed method using the IMCPA in which two outer
therapy arrays transmit 4 MHz CW signals to the target. At the same time, its inner
imaging array emits 6 MHz coded signals similar to conventional sector scanning. The
imaging array receives the reflected coded signals along with reflected therapeutic signals.
After pulse compression, the SNR may be improved and it should be less than -40 dB for
B-mode imaging.
26
Conventional coded excitation employs chirps, Barker codes, and Golay codes
(Golden et al., 1971; Jensen et al., 2005; O’Donnell, 1992). Among them, the Barker
code and chirp signals were chosen for imaging due to its relatively simple hardware
implementation and excellent robustness in noise suppression. The Barker code consists
of N-bit biphase codes, and the optimal peak and range sidelobe level can be obtained
from an autocorrelation function. Its range mainlobe width and sidelobe level depend on
the number of bits and the number of sub-cycles per bit. By using a conventional sidelobe
suppression filter, an acceptable sidelobe level, which is less than -60 dB for B-mode
imaging, can be obtained (Golden at al., 1971). Currently, a 13-bit biphase code sequence
(+1 +1 +1 +1 +1 −1 −1 +1 +1 −1 +1 −1 +1) is the largest length realized for the Barker
code. Figure 2.4(a), (b) shows time and frequency response of the modulated 13-bit
Barker code with 1 cycle per bit. After pulse compression, the range sidelobe level is -
22.3 dB (Figure 2.4(c)) and it is reduced by -53 dB (Figure 2.4(d)) after conventional
sidelobe suppression filtering. In the case of 13-bit Barker code with 2 cycles (Figure
2.5(d)) or 3 cycles per bit (Figure 2.6(d)), they have sidelobe level of around -60 dB after
sidelobe suppression filtering.
27
Figure 2.4: (a) 13-bit Barker code with 1 cycle per bit, (b) frequency response of (a), (c)
after pulse compression, and (d) after conventional sidelobe suppression filtering.
28
Figure 2.5: (a) 13-bit Barker code with 2 cycles per bit, (b) frequency response of (a), (c)
after pulse compression, and (d) after conventional sidelobe suppression filtering.
29
Figure 2.6: (a) 13-bit Barker code with 3 cycles per bit, (b) frequency response of (a), (c)
after pulse compression, and (d) after conventional sidelobe suppression filtering.
The mainlobe in the spectrum of the 13-bit Barker code with 1 cycle per bit goes
beyond the frequency range from 4 MHz to 8 MHz as shown in Figure 2.4(b). This broad
frequency response results in a serious distortion of the mainlobe due to 4 MHz and 8
MHz reflected therapeutic signals. Since more than the 4-cycle-per-bit Barker code might
generate poor axial resolution, 2- and 3-cycle-per-bit Barker codes were considered in
this study. In order to carry out more realistic simulation, transducer impulse response
was used in the program. The -6 dB fractional bandwidths of the imaging and therapeutic
transducer in Figure 2.7 were 50 % and 30 %, respectively. The bandwidth of the
30
therapeutic transducer was lower than that of the imaging transducer due to the lack of
backing of the therapeutic transducer. A 4
th
order Butterworth filter was used to model
transfer functions of these transducers (O’Donnell, 2003).
Figure 2.7: Transducer impulse responses modeled by a 4
th
order Butterworth filter for
point target imaging simulation: (a) Imaging transducer with 50 % -6 dB bandwidth and
(b) therapeutic transducer with 30 % -6 dB bandwidth.
When the 13-bit Barker code with 2 cycles per bit is used, the 4 MHz and 8 MHz
CW interference signals deteriorate the received signal quality for imaging. The
frequency distortion around 4 MHz and 8 MHz arising from the interference signals leads
to a relatively high range sidelobe level, i.e., around -20 dB as shown in Figure 2.8(d).
Although being 16 dB lower than the short pulse signal in Figure. 2.2(d), this level is still
not enough to obtain an acceptable image quality.
31
Figure 2.8: (a) Received 13-bit Barker code with 2 cycles per bit without CW
interference, (b) with CW interference, (c) frequency response of (b), and (d) envelope
signal of (b) after pulse compression with sidelobe suppression filter.
.
The 13-bit Barker code with 3 cycles per bit has about a -50 dB range sidelobe
level in spite of the interference signals as shown in Figure 2.9(d). These two different
results are due to the null point locations in their spectrums. The null points of the 3-
cycle-per-bit Barker code are located around 4 MHz and 8 MHz, thus resulting in
minimized mainlobe distortion.
32
Figure 2.9: (a) Received 13-bit Barker code with 3 cycles per bit without CW
interference, (b) with CW interference, (c) frequency response of (b), and (d) envelope
signal of (b) after pulse compression with sidelobe suppression filter.
The -50 dB range sidelobe level of the 3-cycle-per-bit Barker code is acceptable
for B-mode imaging. However, its axial resolution is poorer than that of the 2-cycle-per-
bit Barker code. The -6 dB axial beamwidths of the original 2- and 3-cycle-per-bit Barker
code in Figure 2.5(d) and Figure 2.6(d) are 0.42 μs and 0.64 μs, respectively. Note that
the -6 dB axial beamwidth of the 2-cycle short pulse signal is 0.34 μs. Therefore, the
focus of this study was on how the range sidelobe level of the 13-bit Barker code with 2
33
cycles per bit could be decreased to at least -40 dB in order to achieve a high axial
resolution.
Another popular coded signal is a chirp signal whose frequency is linearly
changed with time as described below was used as a coded imaging signal:
))
2
)
2
(( 2 sin(
2
0
t
T
B
t
B
f + − π (2.2)
where f
0
is the center frequency and B is the sweeping frequency range, and T is the
duration of the chirp signal. To reduce sidelobe level, several windows have been used
for chirp signals. Among them, the Dolph-Chebyshev window whose ripple ratio and
mainlobe width could be controlled was chosen in this research. Considering the
relationship between the mainlobe width and the sidelobe level of the output signal, -60
dB sidelobe level was defined for the Dolph-Chebyshev window in all simulations and
experiments in this study.
34
Figure 2.10: (a) Time domain response of the Dolph-Chebyshev-windowed chirp signal,
(b) frequency response of (a), and (c) envelope signal of (a) after pulse compression.
Figure 2.10 (a) shows time domain response of the Dolph-Chebyshev-windowed
chirp signal with 20 μm time duration. The frequency sweep range is from 3 MHz to 9
MHz and its sidelobe level was about -68 dB after pulse compression with mismatched
filtering as shown in Figure 2.10 (c).
35
Figure 2.11: (a) Received Dolph-Chebyshev-windowed chirp signal without CW
interference, (b) with CW interference, (c) frequency response of (b), and (d) envelope
signal of (b) after pulse compression with a mismatched filter.
When the HIFU interference signal was mixed with imaging signals, the chirp
signal in Figure 2.11(d) showed -18 dB sidelobe level after pulse compression. Because
the sidelobe level of traditional short pulse signal mixed with HIFU noise was usually
higher than about 2 dB, the chirp signal with only pulse compression had about 16 dB
better than that of normal short pulse excitation. Although the -6 dB axial beam width in
Figure 2.10(d) was 0.45 μs which was 20 % broader than that of the 2-cycle short pulse
36
signal, it might provide reasonable image resolution if the sidelobe level could be reduced
to lower than -40 dB.
2.3.3 Fixed Notch Filtering Technique
Although coded excitation with only pulse compression can decrease the
sidelobe level of HIFU interference about -18 dB, it should be lower than -40 dB for B-
mode imaging. Because the frequency of the HIFU interference is known, notch filtering
technique (Carney 1963) can be used to remove HIFU noise. In general, the frequency
range that contains meaningful information of the image may cover not only in-band
(from 4 MHz to 8 MHz), but also out-band (less than 4 Hz and more than 8 MHz). Thus,
HIFU interference signal should be removed without distortion of in-band and out-band
frequency components of the imaging signal. Fortunately, a reflected CW has a fixed
frequency component, so that the known interference signal may be successfully
minimized with a notch filter capable of rejecting a narrow band of frequency. A notch
filter is widely used in radar or speech processing to attenuate CW signals at specific
frequencies components while minimally affecting other frequency components, and it
can be implemented by relatively simple hardware (Carney, 1963; Hirano et al., 1974;
Waterschoot et al., 2007).
The second order infinite impulse response notch filter based on quadratic
scheme was designed using Matlab and notch attenuation values were found to be around
-37 dB and -31 dB at 4 MHz and 8 MHz, respectively, as shown in Figure. 2.12. Note
37
that the sharpness of the notch filter depends on a quality factor defined as the ratio of
notch frequency over bandwidth of the notch filter (Waterschoot et al., 2007). The quality
factor should be properly determined by considering over shoot or under shoot in a pass
band. In this design, the quality factors for 4 MHz and 8 MHz were 7 and 14, respectively.
The 6 dB attenuation difference between the notch filter responses may be compensated
by the amplitude difference between the fundamental and harmonic components of the
HIFU interference signals. The shallow notch attenuation can maintain original shape but
it generates high remnant ripples behind the mainlobe. Deep notch attenuation can
significantly remove interference signal but it distorts output signal. Because coded signal
has greater robustness in suppressing remnant ripples than normal short pulse signal, to
maintain shape of the original signal as closely as possible, a shallow notch attenuation
value was chosen considering the amplitude of HIFU interference.
38
Figure 2.12: Frequency responses of 4 MHz and 8 MHz notch filters. The notch
attenuation values are -37 dB and -31 dB at 4 MHz and 8 MHz, respectively.
These notch filters were applied to conventional 2-cycle short pulse signal, the
13-bit Barker code with 2 and 3 cycles per bit, and Dolph-Chebyshev-windowed chirp
signal. The amplitude of interference signals was successfully suppressed after notch
filtering in all cases. However, a serious frequency distortion of the short pulse around 4
MHz and 8 MHz generated undesired ripples in its envelope as shown in Figure 2.13(b).
Due to the time response property of the notch filter, the ripples are mainly shown after
the mainlobe.
39
Figure 2.13: Envelopes of imaging signals with CW interference signals: 2-cycle short
pulses, (a) before and (b) after notch filtering; the 13-bit Barker code with 2 cycles per bit,
(c) before and (d) after notch filtering; the 13-bit Barker code with 3 cycles per bit, (e)
before and (f) after notch filtering.
40
Figure 2.14: Envelopes of the Dolph-Chebyshev-windowed chirp signals with CW
interference signals: Before (dashed line) and after (solid line) notch filtering.
In the case of the Barker code with 2 cycles per bit, frequency distortions at
around 4 MHz and 8 MHz were less than the short pulse signal since the locations of null
points of mainlobe were close to 4 MHz and 8 MHz. The Barker code with 2 cycles per
bit as in Figure 2.13(d) had a -40 dB range sidelobe level which was 20 dB lower than the
code without notch filtering, i.e., -20 dB shown in Figure 2.13(c). Since the locations of
first null points of the Barker code with 3 cycles per bit were close to interference
frequencies, its range sidelobe level in Figure 2.13(f) was similar to the pulse
41
compression result without notch filtering (Figure 2.13(e)). This means that the null point
plays a pivotal role in efficaciously decreasing the effect of reflected therapeutic signals
on image quality, and a notch filter helps to further decrease the effect when the null
points do not perfectly match the frequencies of the interference signals. Figure 2.14
shows the windowed chirp signal with CW interference before and after notch filtering.
The sidelobe difference before and after notch filtering was 50 dB in simulation.
Although the small amplitude ripples make mainlobe width broader, as shown at -25 dB
level of the mainlobe, it may be acceptable B-mode imaging given the -40 dB criterion.
2.4 Simulations
2.4.1 Point Target Simulation
To demonstrate the performance of the proposed method by using a novel
transducer, the point target simulation was conducted using a Field-II program which is a
dedicated program for ultrasound field calculation (Jensen, 1996). The 2-cycles short
pulse and the 13-bit Barker coded with 2 and 3 cycles per bit were used as imaging
signals. The design specification of the proposed transducer was summarized in Table 2.2.
The 6 MHz imaging array has 64 elements with 0.6λ = 150 μm pitch, 25 μm kerf, and 8
mm height, while each 4 MHz therapy array has 64 elements with 0.6λ = 225 μm pitch,
25 μm kerf, and 10 mm height. In these pitches, the grating lobe will be shown 74° when
the transducer steer 45° in both transducers. The therapy array with 0.6λ spacing between
42
elements is able to protect normal tissues from the grating lobes of HIFU beam. In the
case of imaging, 74° grating lobe may not be severe problem in real B-mode imaging.
The total dimension of the therapy array was 14.4 mm × 20 mm, which in our
preliminary experiments was sufficient to generate I
SPTA
more than 1000 W/cm
2
at a focal
spot and increase the temperature of soft biological tissue to 70 °C − 90 °C. The whole
dimension of the IMCPA is 14.4 mm × 28 mm. This is proper configuration for
endocavity type transducer. Two 64-element 4 MHz outer row arrays and one 64-element
inner row array were confocally aligned in elevational direction focusing at 40 mm. For
more precise simulation, 0.002 dB/cm·MHz
2
attenuation constant of water was applied to
this simulation.
Table 2.2: Properties for the therapeutic and imaging array.
Property Therapeutic array Imaging array
Center frequency 4 MHz 6 MHz
Wavelength 75 μm 50 μm
Element number 64 (elements) × 2 (rows) 64 (elements)
Pitch/Width/Kerf 225 μm/200 μm/25 μm 150 μm/125 μm/25 μm
Height 10 mm 8 mm
Footprint 14.4 mm × 20 mm 9.6 mm × 8 mm
Focal depth 40 mm 40 mm
F-number (Elevation) 1.4
Intensity (I
SPTA
) 1000 W/cm
2
18.8 mW/cm
2
Grating lobe 74° (steering angle of -45°) 74° (steering angle of -45°)
Figure 2.15 shows geometrical information about apertures with 64 × 3
elements. In Figure 2.15(left), the two outer arrays were activated while inner array was
43
deactivated to transmit therapy signals. For receiving mode, the two outer arrays were
deactivated while inner array was activated to receive reflected therapeutic signal as
shown in Figure 2.15(right). In Figure 2.16, the imaging array was only activated for
transmit/receive mode to collect imaging signals. These two set of data were added
together and used for input data for proposed signal processing method after proper
amplitude compensation using equation (2.1). In the case of receive mode for therapy
array in Figure 2.15(right), reflected therapy signals should be collected by imaging array
with 9.6 mm × 8 mm in Figure 2.16 based on proposed algorithm.
Figure 2.15: Geometrical surface of the therapeutic array aperture combined with a mask
window of the aperture. Outer two arrays are activated for transmit mode (left) and inner
array is activated for receive mode (right).
44
Figure 2.16: Geometrical surface of the imaging array combined with a mask window of
the aperture. Only inner array is activated during both transmit and receive mode.
Using the designed the IMCPA, a point target simulation was performed. The
two outer row arrays transmitted 4 MHz or 8 MHz CW signals and the center-row array
received the reflected interference signals. The center-row array was used to obtain echo
signals for imaging by a transmission/reception process and then the interference signals
were added to the echo signals. Note that the 8 MHz CW signal was regarded as the
second harmonic component of the 4 MHz CW signal. In this simulation, a steering angle
for CW transmission was fixed assuming the following treatment protocol: The CW
beam was focused on a target for a few seconds duration. The -6 dB fractional
bandwidths of the therapy and imaging arrays were same to Figure 2.7. The band stop
attenuation of the notch filter was -37 dB and -31 dB at 4 MHz and 8 MHz, respectively,
as shown in Figure 2.12.
45
Figure 2.17: Point target simulation with imaging signals in which CW interference
signals were superimposed. All figures were logarithmically compressed with a dynamic
range of 40 dB. 2-cycle short pulses, (a) without and (b) with interference; the 13-bit
Barker code with 2 cycles per bit, (c) without and (d) with interference; the 13-bit Barker
code with 3 cycles per bit, (e) without and (f) with interferences.
46
Figure 2.17 shows point target images when therapeutic interference signals were
added to a reflected short pulse signal. All figures were logarithmically compressed with
a dynamic range of 40 dB. The interference signals mixed with a short pulse signal
seriously degraded image quality as presented in Figure 2.17(b). The 13-bit Barker code
with 3 cycles per bit produced a high SNR image (Figure 2.17(f)) because the
interference was greatly suppressed. The 2-cycle-per-bit Barker code generated a high
range sidelobe level that primarily appeared around the center scan line (Figure 2.17(d)).
Figure 2.18 indicates the performance of the notch filter when the interference
was superimposed on the imaging signals. The Barker code with 3 cycles per bit provided
a low noise image shown in Figure 2.18(c) which was similar to the image quality
without the notch filter (Figure 2.17(f)). The image produced by the short pulse signal
shows not only an enhanced range sidelobe level in Figure 2.18(a) due to the notch filter,
but also undesired ripples in the axial direction. Figure 2.18(b) illustrates an improved
range sidelobe level of the 2-cycle-per-bit Barker code compared to that without notch
filtering (Figure 2.17(d)). This point target simulation indicates that a short pulse signal
with a notch filter could be used for B-mode imaging, although there are ripples. In
reality, it might be difficult to completely remove CW signals as demonstrated by the
simulation even using a notch filter. This is because other frequency components around
4 MHz and 8 MHz would be also mixed with the imaging signals. These undesired
interference signals increase the range sidelobe level of the envelope signal of the short
pulse signal, which was experimentally verified.
47
Figure 2.18: Simulated point target images of interference-mixed imaging signals after
notch filtering. All figures were logarithmically compressed with a dynamic range of 40
dB. (a) 2-cycle short pulses, (b) the 13-bit Barker code with 2 cycles per bit, and (c) the
13-bit Barker code with 3 cycles per bit.
48
2.4.2 Intensity Simulation
The pressure and relative/absolute intensity profile simulation using a Field-II
program was performed. The design specification of the proposed transducer was
described in Table 2.2. Figure 2.19 shows a relative intensity contour plot of the transmit
mode for a therapy array when CW signals were used for excitation pulses. Note that this
figure shows transmit sound field considering therapy operation and clinically usable
intensity of the focused HIFU transducer can be obtained from -6 dB DOF (ter Haar,
2000; Azzouz et al., 2006). The -6 dB lateral beamwidth was 1.3 mm and -6 dB DOF was
7 mm. Figure 2.20 shows a relative intensity contour plot of the pulse echo mode for an
imaging array. The -6 dB lateral resolution was 0.7 mm and -6 dB DOF was 16 mm.
Figure 2.21 illustrates the applied pressure value to get required I
SPTA
for therapy, and its
corresponding I
SPTA
is shown in Figure 2.22. The maximum pressure was 5.5 MPa and
I
SPTA
= 1000 W/cm
2
can be obtained from this pressure. Assuming the human tissue target,
0.6 dB/cm·MHz attenuation of human fat was used for this simulation. Figure 2.22 the
3D plot of Figure 2.21 shows spatial intensity distribution, and the maximum intensity at
the center is decreased sharply toward edge. This means that normal tissue outside of the
focal point can be protected during HIFU operation using the IMCPA.
49
Figure 2.19: Relative intensity plot of the therapeutic array: (a) Contour plot, (b)
azimuthal profile, and (c) axial profile.
50
Figure 2.20: Relative intensity plot of the imaging array: (a) Contour plot, (b) azimuthal
profile, and (c) axial profile.
51
Figure 2.21: Pressure plot of the therapeutic array: (a) Axial profile and (b) contour plot.
52
Figure 2.22: Intensity (I
SPTA
) plot for the therapeutic array: (a) Axial profile and (b)
contour plot.
53
Figure 2.23: 3D plot of the intensity (I
SPTA
) for the therapeutic array.
2.5 Prototype Transducer
2.5.1 Modeling
The frequency range for treating malignant prostate tissue is usually between 3 -
4 MHz given the target location, the required intensity level, attenuation, and -6 dB focal
spot size (ter Haar, 2002). In the proposed method, the center frequency of the imaging
transducer is chosen to be between the fundamental and the second harmonic HIFU
54
signal to maximize the available bandwidth for an optimized image. After notch filtering,
not only HIFU noise but also original signal corresponding to notch frequency will be
minimized, so the notch frequencies should be far from the imaging frequency to prevent
from the loss of meaningful image information as much as possible. In this research, 4
MHz and 6 MHz frequencies were chosen for HIFU and imaging, respectively, and thus
imaging transducer had about 67 % maximally available bandwidth between 4 MHz
fundamental and 8 MHz HIFU harmonic signals. This bandwidth became a criterion for
designing transducer and selecting the input waveform. In our prototype transducer, the -
6 dB bandwidth was controlled to be 54 % and the chirp signal with 3 MHz to 9 MHz
frequency sweep range was chosen for imaging signal.
Because the experimental system for multiple channel coded excitation was not
available yet, a prototype transducer composed of three single elements with spherically
focusing was built for a preliminary experiment. However, all design specifications were
based on array type transducer. The 4 MHz therapeutic transducer was designed and its
specification was summarized as shown in Table 2.3. As a piezoelectric material, PZT4
(840, APC International Ltd., Mackeyville, PA) was chosen because it has high
mechanical Q, low dielectric/mechanical loss resulting in reduction of internal loss, and
high Curie temperature to endure increased high temperature. Each element has 1-3
composite structure for pressed focusing process. The aspect ratio which is the ratio rod
thickness over rod width was 0.44, which can reduce side effect of shear wave. The
volume fractional ratio was 79 % for therapeutic transducer and 69 % for imaging
55
transducer. The 79 % volume fraction ratio for treatment was sufficient to generate I
SPTA
than 1000 W/cm
2
at a focal spot. For kerf filler, special epoxy (EPO-TEK314, EPOXY
Technology, Billerica, MA) was used, which is a good epoxy for operation at high
temperature due to it high glass temperature. As a supporting layer, micro-balloon
(Phenolic Microballoons, System Three Ltd., Auburn, WA) whose acoustic impedance is
close to air was used to hold support PZT4. No backing layer was used to maximize the
forward acoustic energy and no matching layer was used for therapy array to minimize
heat absorption during high temperature operation.
Table 2.3: Modeling parameters for the prototype transducer.
Parameters Therapeutic Transducer Imaging Transducer
Center frequency 4 MHz 6 MHz
Piezoelectric material PZT4 PZT-5H
Footprint 14.4 mm × 20 mm 9.6 mm × 8 mm
Thickness 450 μm 280 μm
Rod width/kerf 200 μm /25 μm 125 μm /25 μm
Aspect ratio 0.44 0.45
Volume fraction ratio 79 % (1-3 Composite) 69 % (1-3 Composite)
Focal depth 40 mm 40 mm
Matching layer - EPO-TEK301 (Z=3 MRayl)
Backing layer - WO
3
(4.4 MRayl)
In the case of imaging transducer, regular epoxy (EPO-TEK301, EPOXY
Technology, Billerica, MA) was used for matching layer whose thickness was 110 μm.
As a piezoelectric material, PZT-5H (TFT L-145N, TFT Corporation, Japan) was
selected. The volume fraction ratio was 69 % and its aspect ratio was 0.45. As a backing
layer, tungsten powder loaded-epoxy with impedance of 4.4 MRayl was used. Table 2.4
56
explains the properties of the 1-3 composite piezoelectric materials for imaging and
therapy. The dimension and the acoustic stack are described in Figure 2.24 and Figure
2.25, respectively. This configuration was used for KLM (Krimholtz et al., 1970) model
simulation using tool (PiezoCAD, Sonic Concepts, Woodinville, WA) to predict
electrical impedance and bandwidth of transmit/receive mode.
Table 2.4: Property comparison of the 1-3 composite piezoelectric materials.
PZT4 PZT-5H
Volume fraction ratio [%] 79 69
Density [kg/m
3
] 5808 5739
Longitudinal velocity [m/s] 4126 3843
Coupling coefficient, k
t
0.7 0.67
Relative clamped dielectric constant, ε
33
s
/ ε
0
1655 2119
Acoustic impedance [MRayl] 24 22
57
Figure 2.24: Dimension of (left) therapeutic and (right) imaging transducer. All
transducers are composed of 1-3 piezocomposite.
Figure 2.25: Acoustic stack of (left) therapeutic and (right) imaging transducer.
58
2.5.2 Fabrication
A 14.4 mm × 28 mm prototype integrated HIFU/imaging transducer which
consists of three single elements was fabricated for a preliminary experiment. To make 1-
3 composite for therapeutic transducer, PZT 4 was diced using diamond dicing blade
(Asahi diamond Industry Co., Ltd., Japan) with 25 μm width, and EPO-TEK 314 was
used as a kerf filler. After Gold/Chrome (5000Ǻ/500Ǻ) sputter, lapping process was done
to make thickness of 450 μm. For imaging transducer, PZT-5H and EPO-TEK 301 were
used to make 1-3 composite. Its final thickness was 280 μm after lapping. The surface
conformation process was conducted using a stainless steel fixture. After making two
pieces of therapy elements, they were put in the aluminum housing together. The micro-
balloon was filled as a supporter of therapy elements and fast curing epoxy (Insulcast 501,
ITW Polymer Technologies, Chardon, OH) was used on the surface as a sealant for water
proof. Inside of the housing was divided by using sandwich plates with copper and plastic
layer for RF shielding. 50 Ω coaxial cables were used for interconnection between two
outer therapy elements.
Figure 2.26 (a) shows a spherically focused prototype transducer composed of
three single elements. The front side was covered using Gold/Chrome (5000Ǻ/500Ǻ)
material for HIFU ground while the imaging transducer has different ground (Figure 2.26
(b)) to minimize ground noise comes from therapy elements during high voltage
operation. Thus, the front side of the imaging transducer was not covered by
Gold/Chrome material.
59
Figure 2.26: (a) Photograph of the prototype integrated HIFU/imaging transducer and (b)
cross sectional schematic diagram.
2.5.3 Performance Evaluation
Figure 2.27 and Figure 2.28 show the electrical impedance of integrated
HIFU/imaging transducer by using KLM modeling and impedance analyzer (4294A,
Agilent, USA). The simulated impedance of therapeutic transducer was 10 Ω while the
measured impedance was 21 Ω at 4 MHz. The imaging transducer had 32 Ω of simulation
results and 21 Ω of measured data at 6 MHz. The electromechanical coupling constant k
t
for therapy and imaging transducer were 0.54 and 0.5, respectively.
60
Figure 2.27: Electrical impedance of therapeutic transducer using KLM simulation and
an impedance analyzer: (a) Magnitude and (b) phase.
Figure 2.28: Electrical impedance of imaging transducer using KLM simulation and an
impedance analyzer: (a) Magnitude and (b) phase.
A polished aluminum plate was used to measure the pulse echo signals in the
water tank. The high performance pulser/receiver (5072 PR, Panametrics, USA) was used
to transmit broad bandwidth signal and receive pulse echo data. The received pulse echo
data was acquired using a digital oscilloscope (LC534, LeCroy, Chestnut Ridge, NY). Its
61
frequency spectrum was obtained in Matlab. The measured bandwidth of therapeutic
transducer was 16 % and imaging transducer had 54 %. Their measured center
frequencies were 4.3 MHz and 6.1 MHz, respectively. HIFU transducer has 1170 W/cm
2
of I
SPTA.
Its TAP was 20 W when the loading power was 38 W and thus results in 53 %
transducer efficiency. The TAP of the imaging transducer with 2-cycle PW was 2.5 mW,
mechanical index (MI) was 0.6, and I
SPTA
was 104 mW/cm
2
which satisfy FDA guideline
(Barnett et al., 2000) and will not contribute to any heating on the target.
Figure 2.29: Pulse echo response of the therapeutic transducer in the frequency domain.
62
Figure 2.30: Pulse echo response of the imaging transducer in the frequency domain.
2.6 Experiments
2.6.1 Temperature Profile
The temperature profile for the prototype HIFU transducer was measured by
using a slice of porcine muscle. A freshly excised porcine muscle was obtained from the
butcher. The prototype integrated HIFU/imaging transducer was tested on a sliced
porcine muscle. It was activated during one minute and a thermocouple (TMTSS-020G-6,
63
OMEGA Engineering Inc., Stamford, CT) was placed on the target to monitor
temperature as shown in Figure 2.31. When the spherically focused beam was transmitted
to the target, its temperature increased from 80 °C to 90 °C after 10 seconds with an input
voltage was 140 V
pp
(Figure 2.32(a)). This was a high enough temperature to cause
necrosis of tissues by using the prototype transducer as shown in Figure 2.32(b).
Figure 2.31: (a) Photograph and (b) schematic diagram for temperature measurement
with the prototype integrated HIFU/imaging transducer.
64
Figure 2.32: (a) Measured temperature profile and (b) coagulated sliced porcine muscle
using the prototype integrated HIFU/imaging transducer with spherical focusing.
65
2.6.2 Single Scanline Experiment
Usually, echo signals contain several types of noises that are different from white
noise but can be neglected due to their small amplitude. These noises are associated with
the transducer itself, acoustic loads (Rhyne, 1998), and electronic components (Hayward
et al., 1995). In the case of HIFU, these interference signals might become significant
because of high voltage applied to a HIFU transducer for a rather long duration compared
to the case of imaging. Under this situation, the performances of the proposed method,
evaluated with computer simulation described above, would be degraded. Therefore, the
main purpose of this experiment is to verify whether a coded excitation method with a
notch filter can successfully remove the noises as well as reflected therapeutic signals. A
polished aluminum plate was used to generate strong reflected CW signals including the
noises which were termed spurious signals in this thesis.
Experiments were conducted to obtain imaging signals mixed with CW signals
as described in Figure 2.33. First, 6 MHz imaging transducer was activated by applying
imaging signals to 6 MHz center row of integrated HIFU/imaging transducer. Then, 4
MHz therapeutic transducers were activated by applying CW signals to 4 MHz outer two
rows of integrated HIFU/imaging transducer. Two function generators (33250A, Agilent,
Santa Clara, CA) were utilized to produce both CW and imaging signals such as 2-cycle
short pulses, the 13-bit Barker code with 2 and 3 cycles per bit, and windowed chirp
signal. The transmit signals from the function generator were sent to two different RF
power amplifiers to boost their amplitude and subsequently used to excite the transducers.
66
A RF power amplifier (325LA, ENI Co., Santa Clara, CA) with 50 dB gain was used for
imaging and another amplifier (A300, ENI Co., Santa Clara, CA) with 55 dB gain was
used for generation of therapy signals. With a receiver (5900PR, Panametrics Inc.,
Waltham, MA) and a digital oscilloscope (LC534, LeCroy, Chestnut Ridge, NY), echo
signals were amplified and recorded to perform the proposed signal processing with a
Matlab program. A diode expander (DEX-3, Matec, Northborough, MA) and a diode
limiter (DL-1, Matec, Northborough, MA) were used to protect circuits. As a target, a
polished aluminum plate was immersed into a degassed/deionized water tank and a thin
rubber was placed beneath the aluminum plate to minimize reflected signals coming back
from the bottom of the water tank.
Figure 2.33: (a) Photograph and (b) schematic diagram for a single scanline experiment
to test the performance of the prototype integrated HIFU/imaging transducer.
When the HIFU and imaging transducers were activated simultaneously, high
amplitude 4 MHz and 8 MHz signals were detected by the 6 MHz imaging transducer.
Before notch filtering, one scanline produced by the conventional 2-cycle pulse displayed
67
high amplitude interference signals as shown in Figure 2.34(a). After notch filtering, the
sidelobe level was reduced to less than -30 dB for the 2-cycle pulse case as shown in
Figure 2.34(b). The 13-bit Barker codes with 2 cycles and 3 cycles per bit displayed -40
dB and -44 dB peak sidelobe levels as shown in Figure 2.34(c), (d).
68
Figure 2.34: Measured envelope signals by using the prototype transducer: (a) 2-cycle
short pulse excitation before notch filtering, (b) 2-cycle short pulse excitation after notch
filtering, (c) the 13-bit Barker code excitation with 2 cycles per bit excitation after notch
filtering and pulse compression, (d) the 13-bit Barker code excitation with 3 cycles per
bit after notch filtering and pulse compression. The second peak in the displayed data at
approximately 21 μs is the reflected signal from the bottom of the target.
69
Note that the second large peak at approximately 21 μs in all figures was the
second reflected signal from the bottom of the target. The measured -6 dB axial
beamwidth for the 2-cycle-per-bit code was 18 % broader than 2-cycle short pulse and
34 % narrower than 3-cycle-per bit code. It is speculated that the difference between
simulation and measured data was due to the non-optimized transducer bandwidth and
limited ADC dynamic range. In the case of windowed chirp signal, when the HIFU and
imaging transducers were activated simultaneously, high amplitude 4 MHz and 8 MHz
signals were also detected by the 6 MHz imaging transducer. Before notch filtering, one
scanline produced by the Dolph-Chebyshev-windowed chirp signal with 5 μs time
duration displayed -27 dB after pulse compression as shown in Figure 2.35(dashed line).
After notch filtering, it was decreased lower than -40 dB in Figure 2.35(solid line). Note
that the second large peak at approximately 21 μs in all figures was the second reflected
signal from the bottom of the target. The measured -6 dB axial beamwidth for the 2-
cycle-per-bit code was very close to each other. Note that the mainlobe widths between
two graphs within -40 dB are similar to each other.
70
Figure 2.35: Envelope signals of the Dolph-Chebyshev-windowed chirp signal (dashed
line) before and (solid line) after notch filtering.
2.6.3 In vitro B-mode Experiment
The B-mode imaging experiment was conducted using a 5 cm thick slice of
porcine muscle (in vitro) as the target. In this case, the Dolph-Chebyshev-windowed
chirp signal was used for coded signal. The duration of the signal was 5 μs, frequency
sweeping range was from 3 MHz to 9 MHz, and mismatched filter was used for pulse
compression.
71
Figure 2.36: Photograph with a schematic diagram for single element imaging with the
activated HIFU transducer.
Linear mechanical scanning was used to generate an image because the prototype
transducer was composed of three single elements. Therefore, the HIFU transducer was
also scanned along with the imaging transducer resulting in generation of high amplitude
interference but no coagulated lesions. As shown in Figure 2.36, the 4 MHz HIFU
72
activation equipment was same to Figure 2.33(b) and an imaging system was connected
to the 6 MHz imaging transducer. The imaging system was composed of a linear motor
(LAR37, SMAC Inc., Carlsbad, CA, USA) and a 12-bit ADC (CS12400, Gage Applied
Technologies Inc., Lachine, QC, Canada) controlled by LabVIEW program (LabVIEW,
National instruments Co., Austin, TX). This system was connected to the same
equipment as described in Figure 2.36 to transmit the windowed chirp signal and receive
the pulse echo signals. After obtaining data, post processing such as envelope detection
and logarithmical compression was done using Matlab.
Figure 2.37: B-mode images with a slice of porcine muscle after pulse compression: (a)
Original image of the Dolph-Chebyshev-windowed chirp. After mixing HIFU
interference (b) before and (c) after notch filtering.
Figure 2.37 (a) shows the original porcine muscle image before activating HIFU
transducer and Figure 2.37(b) shows the image after transmitting HIFU beam to the
target. A fixed notch filter with Q values (7/14) was used considering the average
73
amplitude and bandwidth of the interference signals in one frame of image. After notch
filtering, the image in Figure (c) shows clearly that the interference was removed. Note
that some speckle pattern of the recovered image was slightly different from the original
image due to remnant ripples nearby mainlobe, however, this image may be acceptable
for treatment planning and monitoring response of the target in real-time.
2.7 Discussion
In this research, a novel ultrasound transducer design and a scheme to achieve
real-time imaging during treatment were described and the feasibility was demonstrated
through simulation and experimental results. By combining two therapeutic arrays and
one imaging array, the fabrication complexity of the IMCPA may be decreased, and each
array can maintain its own optimal performance. The confocal structure of the IMCPA in
the elevational direction can improve the detection capability. Field-II simulation and
experimental results indicate that coded excitation with a notch filter can improve the
range sidelobe level of the B-mode image during therapy.
In the IMCPA, if high amplitude CW signals generated by therapy arrays are
mixed with a short pulse signal, it would be difficult to extract the imaging signal from
the mixed signal with current signal processing approaches. The coded excitation
technique can overcome this limitation by reducing its range sidelobe level through pulse
compression. In the simulation, when the 13-bit Barker code was used as an imaging
74
signal mixed with CW signals, the Barker code with 3 cycles per bit provided a range
sidelobe level of -50 dB after pulse compression. On the other hand, the 2-cycle-per-bit
Barker code offered a range sidelobe level of -20 dB which was 16 dB lower than that of
a short pulse signal. However, it still could not be used for imaging. It was demonstrated
that this was related to the locations of null points in the frequency response of the Barker
code. The null points of the 3-cycle-per-bit Barker code were located much closer to 4
MHz and 8 MHz frequencies than those of the 2-cycle-per-bit Barker code, thus reducing
the effect of CW interference signals. However, the -6 dB axial beamwidth of the 3-
cycle-per-bit Barker code was 0.3 μs wider than that of the short pulse signal and 0.22 μs
broader than the Barker code with 2 cycles per bit. On the other hand, the Dolph-
Chebyshev-windowed chirp signal had a range sidelobe level of -20 dB, which is similar
to the 13-bit Barker code with 2 cycles per bit.
As a remedy, a notch filter technique was proposed to decrease a range sidelobe
level of the 2-cycle-per-bit Barker code. With the notch filtering, a -40 dB range sidelobe
level of the 2-cycle-per-bit Barker code could be achieved. In the case of the 3-cycle-per-
bit Barker code, its range sidelobe was maintained at about -50 dB level regardless of
whether notch filtering was implemented. Although Field-II simulation in Figure 2.18(a)
shows the feasibility of using a short pulse signal with a notch filter, the experimental
results indicated that the remnant CW components around 4 MHz or 8 MHz after notch
filtering resulted in a -30 dB range sidelobe level. This undesired interference signal may
be reduced by pulse compression of coded excitation successfully. In the single scanline
75
experiment, all coded signals such as 2- and 3- cycle-per-bit Barker code and the Dolph-
Chebyshev-windowed chirp signal had lower than -40 dB sidelobe level.
The frequency sweep range of the chirp signal must be chosen judiciously giving
careful consideration to notch frequencies. If the chirp signal has too wide a frequency
sweep range, the mainlobe width will be narrow but the frequency response containing
important information of the chirp signal will be severely distorted due to HIFU
interference. That is, after notch filtering, main frequency components corresponding to
notch frequencies will be decreased, and thus results in a high sidelobe level. On the
contrary, if the frequency sweep range is too narrow, most frequency information can be
protected from notch frequencies, so the sidelobe level will be considerably reduced but
as a trade off, the mainlobe width which is related to axial resolution will be increased. In
this study, the3 MHz to 9 MHz frequency sweep range was chosen and this condition was
believed to provide proper mainlobe width and sidelobe level. Theoretically, the -6 dB
bandwidth of the imaging transducer should be large enough to include the critical
frequency components of imaging signals but narrow enough to minimally the overlap
with the notch frequencies. In this study, the -6 dB bandwidth for the 6 MHz transducer
was 54 %.
If the notch attenuation value is too shallow comparing to the amplitude of HIFU
interference, the remnant HIFU interference will increase sidelobe level. If it is too deep,
the restored signal will be severely distorted. The coded excitation has low correlation to
CW noise. After just using pulse compression, the sidelobe level could be reduced by
76
about -20 dB. This means that the notch attenuation values of the designed notch filters
can be higher than HIFU interference within the 20 dB margin. These swallow notch
filters combined with pulse compression will allow restoration of images as closely as
possible to the original images. Note that the normal short pulse signal does not have this
notch attenuation margin. If the notch attenuation is shallow than HIFU interference, high
amplitude ripples after mainlobe will be generated.
The pulse compression and notch filtering effect will play an important role in
minimizing HIFU interference to improve image quality. Although real-time lesion
formation could not be observed in the simultaneous mode of imaging and therapy
currently because of system limitation, our experimental results show that if the HIFU
interference could be sufficiently minimized simultaneous B-mode imaging during HIFU
treatment may be possible. In this paper, simulation and experimental results using a
prototype transducer showed that the combination of coded excitation and use of notch
filters on reception is capable of minimizing reflected HIFU signals efficiently and thus
demonstrating the feasibility of the proposed design of an integrated HIFU/imaging
transducer for real-time simultaneous imaging and treatment of prostate tissues.
77
CHAPTER 3: ADAPTIVE SUPPRESSION OF HIFU INTERFERENCE FOR
SIMULTANEOUS THERAPY AND IMAGING
3.1 Introduction
During noninvasive high intensity focused ultrasound (HIFU) surgery, imaging
modalities such as MRI (Magnetic Resonance Imaging) or ultrasound is often used to
allow a physician to have live view of the target being treated (Cline et al., 1995; Ebbini
et al., 2006; Hynynen et al., 1996; Vaezy et al., 2001a). Among them, ultrasound
guidance has been widely used for the treatment of benign and malignant prostate tumors
due to its fast real-time imaging capability and low cost (Azzouz et al., 2006; Blana et al.,
2004; Sanghvi et al., 1999). Typically, a HIFU surgical process is composed of pre-, mid-,
and post- treatment. The physician should identify the location of the target precisely and
determine the proper dose in pre-treatment. In mid-treatment, monitoring the response of
the treated target and tracking the motion of the target or patient are required to ensure
delivering a proper ultrasound dose to the target without damaging normal tissues. An
analysis of the ablated target may be conducted in post treatment (Chinn, 2005).
Especially, by employing B-mode (Brightness mode) ultrasound imaging for monitoring
of the thermal ablation, echogenic regions induced by cavitation of the coagulated lesion
can be monitored during HIFU sonication (Chan et al., 2002; Melodelima et al., 2007;
Owen et al., 2006; Vaezy et al., 2001b).
78
Although ultrasound guidance with an integrated HIFU/imaging transducer can
provide real-time images during pre- and post- treatment, it has had limited success in
realizing real-time B-mode imaging in mid-treatment due to the significant therapeutic
interference received by the imaging transducer (Owen et al., 2006; Vaezy et al., 2001b;
Wu et al., 2008). One of the potential methods to reduce this interference is designing the
imaging transducer at a higher center frequency than that of the therapeutic transducer
(Azzouz et al., 2006; Kluiwstra et al., 1997). Another approach is to transmit mixed-pulse
signals by controlling pulse repetition frequency (PRF) of synchronized imaging and
therapeutic signals (Seip et al., 2002). More recently it was suggested that coded
excitation with the fixed notch filtering (Jeong et al., 2009) might be an alternative for
live imaging during treatment. In this method, the coded signal for imaging and the
continuous wave (CW) signal for therapy were transmitted to the target simultaneously,
and the reflected-mixed signal was received by the imaging transducer. After pulse
compression and fixed notch filtering the final imaging signal had range sidelobe levels
lower than -40 dB while the traditional short-pulse excitation showed higher sidelobe
levels due to the remnant ripples after the mainlobe.
Although coded excitation showed high robustness against residual ripples,
relatively complicated hardware components are required for its implementation, i.e.,
transmission/reception of long pulses per each channel (O’Donnell, 1992). Thus, the
objective of this research is to reduce both reflected therapeutic interference and residual
ripples with traditional short-pulse excitation, accomplished by adaptive noise canceling
79
and subsequently to determine the feasibility of using such a technique for simultaneous
therapy and imaging. Because the proposed algorithm can be implemented with the same
hardware for fixed notch filtering, no additional imaging hardware components will be
needed to produce an acceptable image. A prototype integrated HIFU/imaging transducer
was fabricated for demonstrating the feasibility of the proposed algorithm. All
specifications for the prototype transducer such as the total dimension, center frequencies
for therapy/imaging, and focal depth were designed for satisfying the needs of treating
benign or malignant prostate tumors. Single scanline and B-mode imaging experiments
were conducted with an aluminum plate as a strong reflector and a slice of porcine
muscle, respectively. B-mode image data were collected with a mechanical imaging
scanner during HIFU beam transmission.
3.2 Signal Processing Techniques
3.2.1 Fixed Notch Filtering Technique
In general, the ultrasonic diagnostic image quality during HIFU treatment is
affected by the spurious therapeutic interference. Fortunately, the notch filtering
technique may be used for suppressing this interference because it is sinusoidal with a
known center frequency. A notch filter has been widely used for suppression of
sinusoidal interference in such diverse fields as sonar, speech processing, and
physiological signal processing (Dragosevic and Stankovic, 1995; Ferdjallah and Barr,
80
1994; Hirano et al., 1974). The notch filter can remove specific frequency components
without affecting other frequency components, and it can be implemented by relatively
simplified hardware. Typically, the transfer function of the notch filter can be designed
by finite impulse response (FIR) or infinite impulse response (IIR). An IIR notch filter
can provide a relatively narrower bandwidth than a FIR notch filter and thus produces
less effect on frequencies near the notch frequency. The main challenge however for
designing a fixed notch filter is to determine the proper notch attenuation value against
the interference signals generated by various targets. If the notch attenuation value is
insufficient, high sidelobe level and ripples following the mainlobe of the imaging signal
resulted from remaining interference components will be produced. Since notch
attenuation and bandwidth are directly related, an increase in notch attenuation may add
to the distortion of the filtered imaging signal.
A simulation was carried out to demonstrate this deleterious effect. Here, 4 MHz
and 6 MHz frequencies were used for therapy and imaging, respectively, considering the
location of the prostate tissue is approximately 4 − 5 cm from the rectal surface (Sanghvi
et al 1999). Thus, 4 MHz fundamental and 8 MHz harmonic signals generated by the 4
MHz HIFU transducer severely degrade image quality of the 6 MHz imaging transducer.
Previously, it was shown that two fixed notch filters with 4 MHz and 8 MHz notch
frequencies were required to minimize these interference signals (Jeong et al., 2009).
Figure 3.1 shows the frequency response of the fixed notch filters designed by a 2
nd
order
IIR type based on a quadratic filter with 4 MHz and 8 MHz notch frequencies. When the
81
quality factor Q is 2 (at 4 MHz) and 4 (at 8 MHz), the notch attenuation values (solid
line) are -48 dB and -42 dB, respectively. When Q is 8 (at 4 MHz) and 16 (at 8 MHz), the
notch attenuation values (dashed line) are -36 dB and -30 dB, respectively. In this
simulation, the amplitudes of the fundamental and the second harmonic interference
signals are 40 dB and 34 dB, respectively. Subsequently, the two sets of notch filters so
designed are used to eliminate these interference signals. Thus, one set of the notch filters
has 4 dB shallower notch depth and the other set has 8 dB deeper notch depth than the
amplitude of the interference signal. The 6 dB notch depth difference between 4 MHz
and 8 MHz notch filter responses may be compensated by the amplitude difference
between the fundamental and the second harmonic components of the therapeutic
interference signal.
82
Figure 3.1: Frequency response of the fixed notch filters designed by 2
nd
order IIR type
at 4 MHz and 8 MHz frequency. The double arrow indicates the -3 dB bandwidth of the
notch filter.
83
Figure 3.2: Simulated envelope signals for the fixed notch filters: (a) Original 2-cycle
short-pulse signal. (b) When the therapeutic interference signals are mixed with (a). (c)
After using the fixed notch filter with Q=8/16. (d) After using the fixed notch filter with
Q=2/4.
Figure 3.2 shows the performance of the fixed notch filtering technique. The
sidelobe level of the original 2-cycle short-pulse excitation in Figure 3.2(a) was increased
to approximately -4 dB after being mixed with sinusoidal interference (in Figure 2(b)).
Two types of notch filters (in Figure 3.1) with different Q data were used to remove the
therapeutic interference signal as shown in Figure 3.2(c), (d). One (Q=8/16) of them
shows approximately a -54 dB sidelobe level and high ripples immediately following the
84
mainlobe. The other one (Q=2/4) shows less than -60 dB sidelobe level and lower ripples
than low Q results, but the output signal is distorted significantly when compared to
original signal (Figure 3.2(a)). Therefore, the notch attenuation value must be properly
determined considering the amplitude of the interference signal and the bandwidth of the
imaging signal. However, it is often difficult to select a proper notch attenuation value
using the fixed notch filter as shown in Figure 3.2(c), (d). As a potential solution for this
problem, an adaptive noise canceling method to achieve optimized notch attenuation
values is proposed.
3.2.2 Adaptive Noise Canceling Technique
In HIFU treatment, the signal type and the frequency of the therapeutic
interference are known. It was suggested that adaptive noise canceling technique may
provide the desired signal with optimally cancelled noise components (Glover 1977).
Figure 3.3 shows a block diagram of the proposed adaptive noise canceling algorithm that
may be applied to solving the problem on hand. The primary input signal p(k) is
composed of the original imaging signal s(k) mixed with therapeutic interference n(k).
The reference signal r(k) is the sinusoidal signal with same frequency of the therapeutic
interference signals. This reference signal r(k) will be adaptively matched to the real
interference signal n(k) through an iterative algorithm, and thus the final system output
e(k) will be as close to original signal s(k) as possible. In essence, the adaptive system
algorithm involves an iterative gradient-descent process to find the optimal filter
85
coefficient vectors at the minimum mean-square error position (Widrow 1996). This
LMS (Least-Mean-Square) algorithm can be described by several relative equations.
) ( ) ( ) ( k r k W k y
T
= (3.1)
) ( ) ( ) ( k y k p k e − = (3.2)
) ( ) ( ) ( ) 1 ( k r k e k W k W β + = + (3.3)
where y(k) is the adaptive filter output, W
T
(k) is the N-dimensional transversal filter
coefficient vector for adaptation process, e(k) is the error signal between the primary
input signal and adaptive filter output signal, and β is the adaptation step size. The
βe(k)r(k) is the negative gradient to find the optimal W
T
(k) through the LMS process. As
the adaptive filter output y(k) is approaching n(k), the e(k) will converge to s(k).
Figure 3.4 shows the simulation results of the adaptive noise canceling method
when β is 0.000512 based on same conditions shown in Figure 3.2. The results of this test
are promising as the final output signal shape is very close to original signal. Although,
the -6 dB beamwidth is 11 % broader, the signal pattern as a whole closely matches the
original signal in comparison to the results from the fixed notch filters in Figure 3.2(c)
and 3.2(d).
86
Figure 3.3: Block diagram for (a) the simultaneous therapy and imaging system and (b)
adaptive noise canceling algorithm.
87
Figure 3.4: Simulated envelope signals for the adaptive noise canceling technique (solid
line) and the original signal (dashed line).
3.3 Transducer Design and Fabrication
A prototype integrated HIFU/imaging transducer was fabricated in order to
demonstrate the performance of the proposed algorithm experimentally. This transducer
was composed of three single elements. Inner element was used for imaging and outer
two identical elements were used for therapy. Considering treatment of malignant
88
prostate tissues located at 4 − 5 cm in depth, the dimension of the transducer (14.4 mm ×
28 mm) and the center frequency of HIFU (4 MHz) and imaging transducer (6 MHz)
were chosen.
Figure 3.5: Schematic diagram for the prototype integrated HIFU/imaging transducer
with cylindrically focusing: (a) A photograph and (b) a cross sectional drawing of the
side view.
As a piezoelectric material for HIFU transducer, PZT4 (840, APC Company,
Mackeyville, PA) with a high Curie temperature and a low dielectric/mechanical loss was
selected, and PZT-5H was used for the imaging transducer (Zhang et al., 2005; Zipparo
2003). PZT4 was diced to make 1-3 composite structure by using a 25 μm width blade at
a 250 μm pitch. After cleaning and drying, a high temperature epoxy (EPO-TEK314,
Epoxy Technology, Billerica, MA) was used to fill in the kerfs and cured at 120 ºC for 3
hours. The final thickness of 450 μm was obtained by lapping and a 0.5 μm thick
gold/chrome electrode was sputtered on it. The cylindrically curved shape with 4 cm
89
radius of curvature was obtained through pressing at 130 °C using a stainless steel tube
and a rubber mold. In order to minimize heat absorption and to maximize transmitted
intensity, the prototype transducer did not have a matching or backing layer. The spatial-
peak temporal-average intensity (I
SPTA
) at 4 cm focal point was 350 W/cm
2
and the
transducer efficiency was 54 %.
PZT-5H (TFT L-145N, TFT Corporation, Tokyo, Japan) and unloaded epoxy
(EPO-TEK301, Epoxy Technology, Billerica, MA) were used to make 1-3 composite
structure for the imaging transducer. The fabrication procedure was the same as described
for the therapeutic transducer except that a matching and a backing layer was used. A 110
μm thick layer of unloaded epoxy (EPO-TEK301, Epoxy Technology, Billerica, MA),
was used as the 3 MRayl matching layer, and loaded epoxy (EPO-TEK 301) mixed with
tungsten oxide particles was used for the 4.4 MRayl backing layer. Finally, all
transducers were assembled in the specially designed aluminum housing.
The total acoustic power (TAP) of the imaging transducer with 2-cycle pulsed
wave was 8.5 mW, mechanical index (MI) was 0.7, and I
SPTA
was 144 mW/cm
2
which
satisfy food and drug administration (FDA) guideline (Barnett et al., 2000) and will not
contribute to any heating on the target. The designed transducer has a confocal point at 4
cm to increase detect ability. Because the -6 dB bandwidth was 52 %, the 12 MHz 3rd
harmonic signal could be significantly reduced. Figure 3.5 shows a photograph of the
prototype integrated HIFU/imaging transducer and its cross sectional view. Sheets of
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copper and plastic along with coaxial cables were used inside the housing for radio
frequency (RF) shielding between the HIFU and imaging transducer.
3.4 Experiments
3.4.1 Temperature Profile
The performance of the prototype transducer was tested first to verify whether it
can deliver enough ultrasound energy to heat a target to more than 70 °C. Figure 3.6(a)
shows the setup for this experiment. The tip of a thermocouple (TMTSS-020G-6 and
HH806AU, OMEGA Engineering Inc., Stamford, CT) was placed on the porcine muscle
slice. A high power amplifier (A300, ENI Co., Santa Clara, CA) driven by a function
generator (33250A, Agilent, Santa Clara, CA) was connected to the transducer. The
temperature profile was recorded during 60 seconds. Figure 3.6(b) shows measured
temperature profile, and the temperature of the tissue was increased from 55 °C to 71 °C
during less than 60 seconds.
91
Figure 3.6: (a) Schematic diagram for temperature measurement for the prototype
integrated HIFU/imaging transducer. (b) Measured temperature profile in 60 seconds on a
slice of porcine muscle.
92
3.4.2 Single Scanline Experiment
A strong sinusoidal interference signal was collected from a polished aluminum
plate reflector as shown in Figure 3.7 to verify the effect of the adaptive noise canceling
method precisely. An arbitrary waveform generator (AFG3021, Tektronix, Beaverton,
OR) was connected to the power amplifier (325LA, ENI Co., Santa Clara, CA) with a 50
dB gain. The reflected signal was amplified by a receiver (5900PR, Panametrics Inc.,
Waltham, MA) and recorded by the digital oscilloscope (LC534, LeCroy, Chestnut Ridge,
NY) with 8 bit ADC (Analog to Digital Converter). A diode expander (DEX-3, Matec,
Northborough, MA) and a diode limiter (DL-1, Matec, Northborough, MA) were used to
protect the electrical equipment; same equipment described in Figure 3.6(a) were used to
activate HIFU transducer in this experiment. Note that all measured signals were
obtained using an average-mode of the oscilloscope, i.e. averaging 15 frames per each
signal to minimize white noise interference.
Figure 3.7: (a) Photograph and (b) schematic diagram for a plate experiment for testing
the performance of the prototype integrated HIFU/imaging transducer.
93
Figure 3.8 (a) shows envelope signals of the original 2-cycle short-pulse
excitation when the 8-bit ADC was used. Averaging resulted in a sidelobe level of -60
dB. When the HIFU transducer was activated, the high amplitude interference was added
onto the imaging signal (in Figure 3.8(b)) resulting in -1.3 dB sidelobe level. After notch
filtering with Q=8/16 in Figure 1, the pedestal level was -50 dB but the highest sidelobe
which came from the ripples immediately following the mainlobe was nearly -28 dB in
Figure 3.8(c). With adaptive noise canceling the peak sidelobe level was reduced to -48
dB and the ripples were eliminated (Figure 3.8(d)). A 512-tap FIR band pass filter was
also used in order to remove direct current signal and the third harmonic signal which
was relatively high.
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Figure 3.8: Measured pulse echo envelope signals obtained by the prototype integrated
HIFU/imaging transducer: (a) Original 2-cycle short-pulse excitation. (b) When the
reflected therapeutic interference was mixed with (a) received by imaging transducer. (c)
After notch filtering with a Q being 8 and 16 as shown in Figure 3.1. (d) After adaptive
noise canceling.
3.4.3 In vitro B-mode Experiment
The B-mode imaging experiment was conducted using a 5 cm thick slice of
porcine muscle (in vitro) as the target. Linear mechanical scanning was used to generate
an image because the prototype transducer was composed of three single elements.
95
Therefore the HIFU transducer was also scanned along with the imaging transducer
resulting in generation of high amplitude interference but no coagulated lesions.
Figure 3.9: (a) Photograph and (b) schematic diagram for imaging during HIFU emission
with the prototype integrated HIFU/imaging transducer.
As shown in Figure 3.9, the 4 MHz HIFU activation equipment was same to
Figure 3.6(a) and an imaging system was connected to the 6 MHz imaging transducer.
The imaging system was composed of a linear motor (LAR37, SMAC Inc., Carlsbad, CA,
USA) and a 12-bit ADC (CS12400, Gage Applied Technologies Inc., Lachine, QC,
96
Canada) controlled by LabVIEW program (LabVIEW, National instruments Co., Austin,
TX). This system was connected to the same equipment as described in Figure 7 to
transmit the 2-cycle short-pulse and receive the pulse echo signals. After obtaining data,
post processing such as envelope detection and logarithmical compression was done
using Matlab (The MathWorks Inc., Natick, MA).
Figure 3.10 (a) shows the original porcine muscle image before activating HIFU
transducer and Figure 3.10(b) shows the image after transmitting HIFU beam to the
target. A fixed notch filter with a high Q value (20/40) was used considering the average
amplitude and bandwidth of the interference signals in one frame of image. After notch
filtering, the image in Figure 3.10(c) shows remnant interference components.
Conversely the adaptive notch filtered image, Figure 3.10(d), shows clearly that the
interference was removed. Note that the speckle pattern of the image generated with the
proposed adaptive filtering algorithm is very similar to the original image.
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Figure 3.10: (a) Original image of a slice of porcine muscle, (b) after activating HIFU
transducer, (c) after notch filtering, and (d) after adaptive noise canceling. All figures
were logarithmically compressed with a dynamic range of 40 dB.
98
3.5 Discussion
An adaptive noise canceling algorithm was proposed to effectively suppress
therapeutic interference of varying amplitude while retaining the original signal form as
closely as possible. This algorithm could minimize the reflected therapeutic interference
generated by HIFU transducer and significantly decrease sidelobe ripples immediately
following the mainlobe. The simulation results show that the recovered imaging signal
after applying the proposed algorithm was nearly identical to the original signal.
However the fixed notch filter with too shallow or too deep a notch attenuation value
show high remnant ripples or signal distortion, respectively. The remnant ripple was one
of the main factors that increase the sidelobe level of normal short pulse excitation which
is relatively weaker when compared to the noise signal. Although this artifact may be
reduced with a deeper notch attenuation value, significant frequency distortion around
notch frequency will also result in a distorted signal compared to the original signal.
A prototype integrated HIFU/imaging transducer for treatment of malignant
prostate tissues was fabricated. Although it had a cylindrically focused aperture, it could
increase the target temperature from 55 °C to 71 °C after 60 seconds of operation. This
transducer provided an interference/ripple free single scan-line and B-mode image after
applying adaptive noise canceling. In the presence of a strong reflector, the sidelobe level
after adaptive canceling was -48 dB, whereas the fixed notch filter had -28 dB sidelobe
level. Also, the adaptive noise canceling method could provide a high quality B-mode
99
image with minimized interference and ripples. The stripe pattern artifact in the B-mode
image after fixed notch filtering might come from both remnant interference and ripples
as a result of the fixed notch filtering. This means that fixed notch filtering may not be
able to eliminate the variable interference signal completely while maintaining the
original signal shape. Further, the amplitude of the reflected therapeutic interference may
vary depending on the target property during treatment. The adaptive noise canceling
technique is shown to be capable of overcoming this limitation. The proposed algorithm
may be implemented either in the software or in the hardware with a high speed digital
signal processor (DSP).
It is well known that the B-mode image can be used for observing lesion
formation. However, most ultrasound image guided HIFU systems that implement
simultaneous therapy and imaging suffer from HIFU interference received by imaging
transducer. The proposed method provides a potential solution to minimize this HIFU
interference allowing real-time B-mode imaging during treatment. Although our current
experimental arrangement did not allow monitoring of lesion formation in real-time, the
experimental results demonstrated the proposed algorithm could minimize HIFU
interference. If B-mode imaging is possible during treatment, observation of lesion
formation may be possible based on the hyperechoic response due to cavitation or
different acoustic impedance between a coagulated lesion and a normal tissue.
In this research, it is demonstrated that in traditional short-pulse excitation an
adaptive noise canceling technique could be implemented to reduce therapeutic
100
interference signals and remnant ripples to provide high resolution B-mode imaging
during treatment. Therefore, it may be a promising approach to achieve real-time
simultaneous therapy and imaging in ultrasound image guided HIFU.
101
CHAPTER 4: DUAL-FOCUS HIFU TRANSDUCER FOR EXTENDED
TISSUE LESIONS
4.1 Introduction
In recent years, a number of clinical studies aimed at examining the efficacy of
noninvasive ultrasound therapy have shown that this technique is feasible for treatment of
malignant tissues in organs such as the prostate, liver, kidney, and breast (Blana et al.,
2004; Daum et al., 1999a; Furusawa et al., 2007). Although high intensity focus
ultrasound (HIFU) therapy is effective in destroying tissues non-invasively with shorter
recovery times, and less side effects than invasive surgical techniques (Vaezy et al.,
2001a), it still suffers from a relatively long treatment time for ablation of large tissue
volumes due to a small focal zone per each ultrasound emission (ter Haar, 2000). In the
case of ultrasound therapy for the malignant prostate tissue, a target is virtually divided
into anterior, middle, and posterior zones, and each zone is separately treated. Thus, the
overall treatment time may be as long as several hours depending on a target size (Chinn,
2005).
Several techniques have been developed to decrease the treatment time of
noninvasive ultrasound therapy. One of them is the generation of multi-foci with a phased
array transducer capable of electronic focusing and steering (Daum et al., 1999b; Ebbini
102
et al., 1989; Wan H et al., 1996). Also, multi-zone transmit focusing which is usually
used for imaging can increase ablated tissue lesions in the axial direction (Do-Huu and
Hartemann, 1982). However, the conventional multi-foci technique and multi-zone
transmit focusing scheme requires a complex array architecture. Also, numerous
electronic components including high-power amplifiers and time delay components are
required to activate and control each element independently. Although several
investigators have used specially designed lenses, and conical shape transducer to
increase depth-of-focus (DOF), most studies were focused on enhancing imaging
performance (Burckhardt et al., 1973; Patterson and Foster, 1982; Trzaskos and Young,
1987). A few researchers have proposed a split-focusing technique to generate multi-foci
simultaneously with a geometrically divided transducer, or a transducer with sectional
electrode (Patel et al., 2008; Sasaki et al., 2003; Seip et al., 2001) driven by voltages of
different phases. The aforementioned studies have focused on producing broad tissue
lesions in the lateral and elevational directions.
This research focuses the development of a specially designed therapeutic
ultrasound transducer for reducing the treatment time of large tumors in high intensity
ultrasound therapy by increasing DOF in the axial direction. The proposed transducer is
composed of a concentric disc- and an annular-type element and each element has a
different radius of curvature to produce two focal zones upon one excitation called dual-
focus therapeutic ultrasound transducer (DFTUT). Two different center frequencies are
used for simultaneous dual-zone transmit focusing and thus result in an enhanced focal
103
depth. Each element is made of a piezoelectric composite material of a different thickness
and subsequently optimized for transmitting ultrasound corresponding to its own resonant
frequency. Thus, both elements work like band-pass filters of the excitation signal. These
properties enable the proposed transducer to be activated by a single transmitter capable
of generating a dual-frequency mixed signal. The center frequency and the dimension of
each element can be optimized based on the application considering a target size and
distance. Also, the relative geometric focus offset between two different focal points and
the output power of each element may affect the uniformity of the extended DOF and the
lateral beamwidth. A good alignment between these two elements in the fabrication
process is a critical step to achieve a uniform compound beam profile in the axial and
lateral direction.
In this study, a prototype DFTUT was built and the performance was
demonstrated by hydrophone measurements and lesion formation tested on a piece of
beef liver. Sound field simulations were conducted to obtain proper design specifications
for the DFTUT and predict its performance. Also, a bio-heat transfer simulation showed
the estimated temperature distribution in the extended DOF for the DFTUT.
104
4.2 Methods
4.2.1 Transducer Design
Clinically the efficacy of a focused therapeutic transducer may be estimated from
its -6 dB intensity contour of the focal zone (ter Haar, 2000). Along the axis of
propagation, the effective focal zone is defined by the DOF which is related to the square
of the f-number (focal depth/aperture size) and the wavelength (Shung, 2006).
Unfortunately the DOF is sometimes sacrificed for the sake of generating sufficiently
high intensity for tissue treatment. Thus, a specially designed DFTUT is proposed as
shown in Figure 4.1 to increase the DOF while maintaining the necessary ultrasound
intensity level for therapy. The DFTUT is composed of a disc-type-inner element and an
annular-type-outer element with different radii of curvatures. The number of annular
elements can be increased depending on the desired DOF. In this thesis, a single annular
element was used for simulation and the preliminary experiment. Malignant prostate
tissue was chosen as a specific therapeutic target for this transducer.
105
Figure 4.1: Schematic diagram of the dual-focus therapeutic ultrasound transducer
(DFTUT) with specification: (a) Side view and (b) front view. Note that the relative
geometric focus offset between two focal points is 5.24 mm.
106
Therapeutic ultrasound signals in the 3 – 4 MHz range are typically used for
noninvasive surgery of the malignant prostate tissue, and capable of producing high
thermal damage with minimal cavitation (Blana, 2004; Sanghvi, 1999). Therefore, the
inner and outer elements of the DFTUT have 4.1 MHz and 2.7 MHz center frequencies
with 19 mm and 24 mm radii of curvatures, respectively. In this case the relative
geometric focus offset between two focal depths was 5.24 mm as shown in Figure 4.1(a).
The relative geometric focus offset was determined considering the overlapped zone of
two elements at higher than -6 dB. The electrodes of the these elements were connected
together allowing ultrasound at two different frequencies to be simultaneously emitted to
the two focal zones with a single transmitter which generates a continuous wave (CW)
signal at two-frequencies. The relative intensity for each element can be controlled by
changing the amplitude of the two frequency components in the input signal. The
electrical impedance for inner and outer elements should be matched to that of the system
mainly with a power amplifier for balanced output power.
The diameters of the inner and outer elements were 12 mm and 21 mm,
respectively. Under these conditions, the -6 dB lateral beamwidth of the DFTUT was
similar to those obtained by inner and outer elements excited independently. The overall
DOF for the DFTUT can be increased by minimizing the overlapping DOF regions
generated by these two elements.
107
4.2.2 Sound Field Simulation
To evaluate the performance and to obtain design specification of the DFTUT,
the on-axis transmit beam profile was computed with the Field-II (Jensen and Svendaenl,
1991) simulation program derived from the Tupholme-Stepanishen method (Stepanishen,
1971; Tupholme, 1969). The two signals were simultaneously emitted to the target via
the inner and outer elements of the dual curved aperture in order to model the amplitude
and phase interaction between two waves of different frequencies in the transmit-field.
The dual curved aperture was created with Field-II’s automatic meshing aperture
generation as shown in Figure 4.2, and two waveforms of different frequencies were
assigned to each element. This method was compared to post-sum approach, i.e., sum of
two transmit-field profiles for each aperture: There was no significant discrepancy in the
time and frequency responses of scanlines and final beam profiles implemented by these
two schemes.
The transducer aperture was apodized using the Hanning window and segmented
by 300 μm rectangular elements. The radii of curvatures for inner and outer elements are
19 mm and 24 mm, respectively. The 4.1 MHz and 2.7 MHz input signals of 30 % -6 dB
bandwidth were assigned to each aperture based on the expected output from the
prototype transducer without a backing layer. The sampling frequency was 40 times of
the 4.1 MHz frequency, and the amplitude of each frequency component was the same.
The lateral and axial grid sizes were 20 μm and 100 μm, respectively. The relative
geometric focus offset between two focal points was 5.24 mm. In this distance, the -2 dB
108
contours off the maximal peaks for the two signals in Figure 4.3 and Figure 4.4 cross
each other.
Figure 4.2: Schematic diagram of the DFTUT aperture used in the Field-II simulation.
The radii of curvatures for inner and outer elements are 19 mm and 24 mm, respectively.
The diameters of inner and outer elements are 12 mm and 21 mm, respectively.
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Table 4.1: Specification of the DFTUT for sound field simulation.
Inner Element Outer Element
Center frequency [MHz] 4.1 2.7
Diameter [mm] 12 21
Geometric focus [mm] 19 24
F-number 1.6 1.1
Relative geometric focus offset [mm] 0 5.64
*Water attenuation coefficient [dB/cm·MHz
2
] 0.0025
*Fat attenuation coefficient [dB/cm·Hz] 0.6
* (Shung, 2006)
Table 4.1 shows the specification of the DFTUT. Simulations were conducted
using the properties of the two different media such as water and fat. The results of these
simulations are summarized in Table 4.2 and are displayed for water only in Figure 4.3 −
Figure 4.5. Figure 4.3 and Figure 4.4 show the transmit beam profiles of the DFTUT’s
inner and outer elements excited independently. The contour plots in Figure 4.3(a) and
Figure 4.4(a) display the relative intensity in the decibel scale. Figure 4.3(b) and Figure
4.4(b) show the lateral beam profile of each element at maximal peak. The -6 dB lateral
beamwidths of inner and outer elements were 0.81 mm and 0.72 mm, but the sidelobes
were -17 dB and -9 dB. Because outer element has a ring type aperture, the sidelobe is
higher and the lateral beamwidth is narrower than a disk type aperture. Figure 4.3(c) and
Figure 4.4(c) show the axial beam profile of each element, and -6 dB DOF were 9.1 mm
and 11.7 mm, respectively.
Figure 4.5 shows transmit beam profile of the DFTUT at water medium. Its -6
dB DOF was 1.7 and 1.3 times larger than those obtained with inner and outer elements,
110
respectively. The -6 dB lateral beamwidth of the proposed transducer was 0.75 mm
which was close to those of inner and outer elements excited independently. The -9 dB
sidelobe level was close to outer element. As a comparison, a single focused transducer
with design parameters equal to the mean of the design parameters for the DFTUT
(center frequency 3.4 MHz, focal depth = 21.5 mm and diameter = 21 mm) was also
simulated (Figure 4.6). The DFTUT’s -6 dB DOF was 3.5 times larger, its -6 dB lateral
beamwidth was 0.11 mm broader, and the sidelobe was 9 dB higher than those obtained
with the single focused transducer. In the case of fat, the extended -6 dB DOF for the
DFTUT was 3.4 times larger than 3.4 MHz single element transducer. The sidelobe of the
DFTUT can be changed by controlling the dimension of outer element. Typically, the
high sidelobe level may cause undesired tissue damage in adjacent normal tissues.
However, sometimes it was used for the broader lesion formation to decrease treatment
time of a large tumor (Wu et al., 1999).
111
Figure 4.3: Transmit beam profile of the inner element with 19 mm focal depth: (a) A
contour plot in the decibel scale, (b) lateral beam profile, and (c) axial beam profile.
112
Figure 4.4: Transmit beam profile of the outer element with 24 mm focal depth: (a) A
contour plot in the decibel scale, (b) lateral beam profile, and (c) axial beam profile.
113
Figure 4.5: Transmit beam profile of the DFTUT: (a) A contour plot in the decibel and
(b) axial beam profile.
114
Figure 4.6: Transmit beam profile of the single focused transducer with 21.5 mm focal
depth: (a) A contour plot in decibel, (b) lateral beam profile, and (c) axial beam profile.
115
Table 4.2: Simulated -6 dB DOF, -6 dB lateral beamwidth, and sidelobe of the
transducers.
Medium
Water Fat
4.1 MHz Inner
Element
-6 dB DOF [mm] 9.1 8.5
-6 dB Lateral
beamwidth [mm]
0.81 0.8
Sidelobe [dB] -17 -17
2.7 MHz Outer
Element
-6 dB DOF [mm] 11.7 11.1
-6 dB Lateral
beamwidth [mm]
0.72 0.71
Sidelobe [dB] -9 -9
3.4 MHz Single
Element
-6 dB DOF [mm] 4.4 4.3
-6 dB Lateral
beamwidth [mm]
0.64 0.63
Sidelobe [dB] -18 -18
DFTUT
-6 dB DOF [mm] 15.4 14.6
-6 dB Lateral
beamwidth [mm]
0.75 0.73
Sidelobe [dB] -9 -9
4.2.3 Transducer Fabrication
A prototype DFTUT was built following the specifications summarized in Table
4.3. This transducer has 1-3 piezoelectric composite elements using PZT4 (840, APC
Company, Mackeyville, PA) and epoxy (EPO-TEK314, Epoxy Technology, Billerica,
MA) so that they can be spherically shaped, and to reduce acoustic impedance mismatch
between a transducer and a medium. PZT4 has high Curie temperature and high
mechanical Q and thus is an excellent material for therapeutic transducer designs. Also,
the epoxy used has a high glass transition temperature (100 °C) which makes it less
susceptible to failure during operation. For the 4.1 MHz inner element, a pre-poled PZT4
116
plate was diced with a 35 μm wide blade to make the 1-3 composite. The composite pitch
was 250 μm and the post-widths were 215 μm. After a cleaning and drying procedure the
kerfs were filled with unloaded EPO-TEK 314. After the composite was cured at 120 °C
for 3 hours, the composite was then lapped to a final thickness of 450 μm and heat
pressed to a spherically curved shape at 130 °C using a rubber mold and a chrome/steel
ball. The same process was used to fabricate the outer ring element. After heat pressing,
the two elements were carefully bonded together, and a 0.5 μm thick gold/chrome-
sputtered layer was used as the electrode for both the signal and ground electrodes. No
matching and backing layers were used in the prototype transducer in order to minimize
heat absorption and to maximize transmit intensity. Figure 4.7 shows a photograph of the
finished prototype DFTUT.
Table 4.3: Parameters for fabrication of the prototype DFTUT.
Inner Element Outer Element
Piezoelectric material PZT4 PZT4
Epoxy EPO-TEK 314 EPO-TEK 314
Center frequency [MHz] 4.1 2.7
Inner diameter − 12
Outer diameter [mm] 12 21
Composite pitch [μm] 250 250
Composite kerf [μm] 35 35
Thickness [μm] 450 600
Volume fractional ratio [%] 74 74
Post aspect ratio (width/thickness) 0.48 0.36
Focal depth [mm] 19 24
117
Figure 4.7: Photograph of the prototype DFTUT.
4.3 Experiments
4.3.1 Electrical Impedance Measurement
Figure 4.8 shows the measured electrical impedance of the water-loaded DFTUT
with an impedance analyzer (4294A Impedance Analyzer, Agilent, Santa Clara, CA).
Two peaks of the impedance plots are seen in series. One is for the outer element at a
lower frequency and the other for the inner element at a higher frequency. The anti-
resonance frequencies for each element were 42 Ω at 2.7 MHz and 93 Ω at 4.2 MHz,
respectively. The driving frequencies of the DFTUT were determined given a maximal
peak of the transmit response resulting in a peak pressure value in the hydrophone
118
measurement. The 2.7 MHz and 4.1 MHz frequencies for outer and inner elements
yielded the highest pressures, and their impedances were 40 Ω and 60 Ω, respectively.
Figure 4.8: Measured electrical impedance of the DFTUT with a water load. There are
two impedance peaks in series for inner and outer elements.
4.3.2 Transmit Response Measurement
A needle hydrophone (HPM04/1, Precision Acoustics Ltd, Dorchester, UK) was
used to measure the transmit response of the DFTUT as shown in Figure 4. 9. A function
generator (33250A, Agilent, Santa Clara, CA) capable of generating a 2-cycle PW signal
at frequencies at 4.1 MHz and 2.7 MHz was connected to a 50 dB RF power amplifier
(325LA, ENI Co., Santa Clara, CA) resulting in 32 V
pp
input voltage and subsequently
used to activate the DFTUT. The distance between the DFTUT and the hydrophone was
119
varied from 5 mm to 40 mm via a controller (6000ULN, Burleigh Instruments Inc.,
Fishers, NY) with a XYZ translation stage driven by a piezoelectric stepper motor (IW-
700 Series Inchworm Motor, Burleigh Instruments Inc., Fishers, NY). Signals received
by the hydrophone were amplified by 25 dB (Hydrophone Booster Amplifier, Precision
acoustics LTD., UK), measured with a digital oscilloscope (LC534, LeCroy, Chestnut
Ridge, NY) with 8-bit ADC (Analog to Digital Converter) card, and recorded by a
computer with a data acquisition board.
Figure 4.9: Experimental setup for measurement of the transmit response, DOF, and
lateral beamwidth of the DFTUT by using a hydrophone: (a) A photograph and (b) a
schematic diagram.
120
Figure 4.10 shows the measured transmit frequency domain response for the
DFTUT at various depths along the axial direction. Its magnitude at different depths was
observed to be proportional to the amount of energy contributed by the two elements. In
the near field as shown in Figure 4.10(a), 4.1 MHz frequency component of the inner
element was higher than 2.7 MHz frequency, and this ratio was reversed at far field as
shown in Figure 4.10(b), (c), (d). The distances between a hydrophone and the transducer
in Figure 4.10 (a) − (d) were 18 mm, 23 mm, 28 mm, and 33 mm, respectively.
Figure 4.10: Frequency domain plots of the measured transmit response along the axial
direction: (a) 18 mm, (b) 23 mm, (c) 28 mm, and (d) 33 mm in depth.
121
4.3.3 DOF/Lateral Beamwidth/Sidelobe Measurement
The same experimental setup as described in Figure 4.9 was used to measure the
-6 dB DOF and the -6 dB lateral beamwidth of the DFTUT. The swept range for the
hydrophone was from 5 mm to 40 mm in the axial direction and from -5 mm to 5 mm in
the lateral direction. Eight sets of lateral beam profiles were recorded at 2 mm intervals
within the -6 dB DOF, i.e., 15 mm − 30 mm. The measured -6 dB DOF was
approximately 14.5 mm as shown in Figure 4.11(a). Figure 4.11(b) shows measured -6
dB lateral beamwidths about 15 mm − 30 mm in the -6 dB DOF, however, a dip was
observed around 25 mm.
There are several possibilities that may explain this phenomenon including a
difference in the delivered energies between the near field and far field and misalignment
between two elements during the fabrication process, error of radius of each curvature,
and the offset between two elements during pressed focusing. Because the most likely
cause may be resulted from the difference in delivered energies due to the impedance and
dimension of each element, a modified simulation was conducted by reducing the
amplitude of the outer element signal to half of the inner element considering the
amplitude of the dip in Figure 4.11(a), also including geometrical errors of the prototype
transducer such as dimension and the depth offset between two elements. However, it
appeared that the most significant change came from the amplitude difference between
the two frequencies. When the amplitude of low frequency component was decreased to
half of the high frequency component in the input signal, the resulted curve was
122
remarkably similar as shown in Figure 4.11(a), (b). Although the location of the dip point
is lower than measured data, the patterns were very similar. Thus, compared to same
amplitude results in Figure 4.5, it is clear that the output power of each element may
affect the performance of the DFTUT. It seems that this phenomenon might be caused by
the electrical impedance mismatch due to dimension and driving frequency for each
element.
The alignment accuracy may be also critical during the fabrication process.
These results showed that the extended -6 dB DOF could be obtained although it was not
possible to maintain a constant lateral beamwidth throughout the DOF which likely was
caused by the difference in the delivered energies as illustrated in Figure 4.11(a). This
unbalance in emitted power between the two elements may be minimized through more
optimized transducer design carefully considering the impedance match, frequency,
dimension of each element. Figure 4.12 shows the measured lateral beam profiles at 20
mm and 22 mm depths with simulation data under the condition which is similar to
Figure 4.11. The measured and the simulated sidelobe level were -11.3 dB and -13.4 dB,
respectively. The discrepancies between the two data sets might be attributed to the
limited dynamic range of the 8-bit ADC card.
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Figure 4.11: Simulated and measured data for the DFTUT using a hydrophone: (a) An
axial beam profile with DOF and (b) -6 dB overall lateral beamwidth within the -6 dB
DOF.
124
Figure 4.12: Simulated and measured lateral beam pattern for the DFTUT: The
simulated (red-solid line) and the measured data at 20 mm (black-solid line) and at 22
mm in depth (blue-dashed line). Note that the pedestal level of the all measured data is
higher than simulated data due to ADC noise.
4.3.4 Bio-Heat Transfer Simulation and In vitro Experiment
A bio-heat transfer simulation was performed with the simulated acoustic
pressure field in which the amplitude for outer element was lower than inner element to
obtain more realistic simulation results as shown in Figure 4.11. The Pennes bio-heat
125
transfer equation (Pennes 1948) was approximated by the equation below and
numerically solved with Matlab (The MathWorks Inc., Natick, MA).
q T T c W T k
t
T
c
a b b t t t
+ − + ∇ =
∂
∂
) (
2
ρ (4.1)
where
t
ρ is the tissue density, C
t
is the specific heat of tissue, k
t
is the tissue thermal
conductivity, W
b
is the blood perfusion rate, C
b
is the specific heat of blood, T
a
is the
arterial temperature, T is the tissue temperature, and q is the absorbed ultrasound power
density defined below (Nyborg, 1981). Table 4.4 shows the parameters used for this
simulation (Diederich and Burdette, 1996) targeting a soft tissue.
Table 4.4: Parameters for bio-heat transfer simulation.
Density, ρ
t
[kg/m
3
] 1050
Velocity of sound, c [m/s] 1540
Specific heat for tissue, c
t
[J/kg/ºC] 3639
Specific heat for blood, c
b
[J/kg/ºC] 3825
Blood perfusion, W
b
[kg/m
3
/s] 5
Thermal conductivity, k
t
[W/m/ºC] 0.56
Ultrasound absorption coefficient, α [Np/m/MHz] 41
Time step [sec] 0.005
Simulation time [sec] 30
c
p
q
ρ
α
2
= (4.2)
where α is the acoustic absorption coefficient and p is the measured pressure at focal
point, ρ is the density and c is the velocity. A numerical finite-difference method was
126
used for solving the bio-heat transfer equation by replacing the derivative equation with
difference quotients. The X axis range was from -4 mm to 4 mm and the Z axis range was
from 1.5 mm to 46.5 mm. The X- and Z- step sizes were 0.02 mm and 0.4 mm,
respectively. The time step was 0.005 seconds and simulation time was 30 seconds.
Figure 4.13 shows the estimated temperature distribution for the DFTUT in 3D modeling.
Note that in this figure the position of the transducer is on the left side as indicated by a
depth of zero. The maximum temperature recorded was 69 ºC at 20 mm which is near the
geometrical focus of the inner element, and the second peak has 52 ºC at 26 mm which is
near the geometrical focus of the outer element.
A test on soft biological tissue lesion formation was conducted to verify the
performance of the device. A test on soft biological tissue lesion formation was
conducted to verify the performance of the device. A freshly excised beef liver was
obtained from the butcher. The experimental arrangement was nearly identical to what
was described for the hydrophone test as shown in Figure 4.14(a). The target was placed
in a water bath and the target distance from the surface of the transducer was controlled
to be approximately 10 mm. The amplitude of the signal used to excite the transducer was
increased to 140 V
pp
with a 55 dB power amplifier (A300, ENI Co., Santa Clara, CA)
resulting in an estimated 6.1 MPa peak positive pressure. Figure 4.14(b) shows the
ablated tissue lesion of a beef liver with 140 V
pp
applied voltage for 30 seconds. The
lesion size was approximately 20 mm in length and tapering width from about 8 mm.
This tapering was likely due to the difference in the delivered energies between the two
127
elements in the DFTUT. The coagulated region in the axial and lateral direction was
wider than the results obtained by hydrophone measurements in Figure 4.11, which may
be related to the thermal conduction due to high ambient temperature based on extended
exposure time about 30 seconds (ter Haar et al., 1989). Note that the coagulated lesion in
Figure 4.14 was formed in front of the geometrical focal depths of the DFTUT and
moved toward the surface of the tissue. This phenomenon may be explained by the
nonlinear distortion of the ultrasound wave in the tissue. Another reason may be the
changed property of the ablated-tissue resulting in abnormal attenuation coefficient
during high temperature HIFU sonication (Chen et al., 1997; ter Haar et al., 1989).
128
Figure 4.13: Simulated temperature distribution for the DFTUT. The position of the
transducer was on the left side as indicated by a depth of zero.
129
Figure 4.14: Cross-section of a piece of beef liver after HIFU sonication with the
DFTUT. The arrow indicates that the HIFU exposure direction. (a) A schematic diagram
to generate (b) the ablated lesion in a beef liver specimen by using the DFTUT.
4.4 Discussion
In this research, a proposed method can increase the volume of tissue lesion in
the axial direction by extending the DOF with the DFTUT targeting the treatment of
prostate tumors. The basic concept of this method is partially overlapping two DOFs
generated by spatially aligned two coaxial elements, which have different center
frequencies. To extend the DOF maintaining the uniformity in the axial and the lateral
130
beam profiles, transducer design must be optimized considering the frequency, the
dimension of each element, and the relative geometric focus offset. In the sound field
simulation, the 4.1 MHz and 2.7 MHz center frequencies for the inner and outer elements
and 5.24 mm of the relative geometric focus offset between two focal depths could
provide extended -6 dB DOF about 3.4 − 3.5 times larger than a conventional transducer
of the same diameter while maintaining a relatively uniform lateral beamwidth.
The proper relative geometric focus offset can be determined by the cross point
of two intensity plots at which the intensity sum of two overlapped zones belonging to
inner and outer elements should be higher than -6 dB. However, in reality, because the
slope of each curve in the overlapped zone is not symmetric, a higher value may be
required. In this study, -2 dB point from the maximal intensity was chosen and thus two
focal depths were separated by 5.24 mm. These results demonstrate that the length of the
extended DOF can be changed by controlling the relative geometric focus offset. The
lateral beamwidth was similar to other transducers as shown in Table 4.2 because this
method is mainly focused on the extended DOF in axial direction. The sidelobe level of
the DFTUT also can be varied by changing the physical dimensions of the two elements.
In this study, an approach in transducer fabrication in which two elements were
excited simultaneously with a single electrode by a single transmitter instead of two
separated electrodes was introduced to extend the DOF. The prototype transducer based
on this design could provide 23 mm extended DOF resulting in a longer lesion. In this
configuration, electrical impedance matching for both elements and the total input
131
impedance may be important for balancing the output power. The length of the extended
-6 dB DOF with a hydrophone measurement shows 14.5 mm although there was a dip
around 25 mm. One of the possible reasons may be the difference in delivered energies
between these two elements due to non-optimized electrical impedance of the prototype
transducer. In sound field simulation, when the amplitude of the low frequency is lower
than the high frequency components, the final output DOF shows a trend similar to the
measured data with a nonuniform intensity contour. This observation seems to suggest
that output power unbalance may the primary culprit of this behavior, and it may be
solved by better matching electrical impedances or by changing the excitation parameters.
Several fabrication issues such as a depth offset and alignment errors during the
assembly of the two elements can be solved with the use of proper fixtures. Additionally,
under similar driving conditions the ultrasound intensity along the elongated DOF
produced by the DFTUT may be lower than a single focused transducer, so higher
driving power may be required to obtain the same treatment effect. The bio-heat transfer
simulation results show that the DFTUT can generate a temperature up to 69 ºC under the
current driving conditions within the extended DOF and the lesion formation test with a
piece of beef liver shows a coagulated lesion of 23 mm length and 8 mm width with a
tapering shape, which may come from the relatively lower intensity of the outer element.
The size of the coagulated lesion for in vitro experiments was larger than hydrophone
measurement data. One of the potential reasons causing this phenomenon was the thermal
conduction related to an extended sonication time. The changed attenuation coefficient
132
and nonlinear distortion of the ultrasound wave in the ablated lesion during HIFU
sonication may cause formation of the lesion in front of the geometrical focus.
These results demonstrate the feasibility of the DFTUT to extend DOF resulting
in a broad tissue lesion in the axial direction. This technique may be useful for treatment
of the large tumors especially deep-seated tumors. The fabrication process for the
DFTUT may be further improved to optimize the performance of such a device.
133
CHAPTER 5: SUMMARY AND FURTHER WORKS
5.1 Summary
For noninvasive treatment of the prostate tissue using HIFU, this thesis proposes
a design of the integrated HIFU/imaging transducer called IMCPA and signal processing
techniques to perform both therapy and imaging simultaneously. The proposed integrated
HIFU/imaging transducer is composed of three rows of phased array elements confocally
aligned in the elevational direction. The center row forms the 64-element 6 MHz imaging
array and the two identical 64-element 4 MHz outer rows work together to produce the
HIFU signal. Different stack configurations and piezoelectric materials were used to
achieve the desired performance as a combination HIFU and imaging transducer. The
total dimension of the HIFU array was 14.4 mm × 20 mm, which in the preliminary
experiments was sufficient to generate I
SPTA
more than 1000 W/cm
2
and to increase
temperature more than 90 °C at a focal spot. The 13-bit Barker code with 2 and 3 cycles
per bit, and the Dolph-Chebyshev-windowed chirp signal were used for implementing
coded excitation. Subsequently, the second order infinite impulse response notch filters
were used to remove the remnant HIFU interference composed of 4 MHz fundamental
and 8 MHz harmonic signals from the 6 MHz imaging echo response.
134
A 14.4 mm × 28 mm prototype integrated HIFU/imaging transducer which
consisted of three single elements with spherically focusing was fabricated as a
preliminary experiment. When the HIFU and imaging transducers were activated
simultaneously, high amplitude 4 MHz and 8 MHz signals were detected by the 6 MHz
imaging transducer. After notch filtering, one scanline produced by the 13-bit Barker
code with 2 cycles or 3 cycles per bit and the Dolph-Chebyshev-windowed chirp signal
displayed a sidelobe level lower than -40 dB while conventional 2-cycle pulse had a
maximum sidelobe level of -30 dB. These results agree with those obtained with a Field-
II simulation. The B-mode image with a slice of porcine muscle showed the recovered
image was very similar to the original image. These experimental results prove that the
combination of coded excitation and use of notch filters on reception can minimize
reflected HIFU signals efficaciously, and thus the proposed design of an integrated
HIFU/imaging transducer can be used for real-time simultaneous imaging and treatment
of the prostate tissue.
The shortcoming is, however, that relatively complicated electronics may be
needed to utilize coded excitation in an array imaging system. It is for this reason that in
this research an adaptive noise canceling technique was proposed to improve image
quality by minimizing not only the therapeutic interference, but also the remnant sidelobe
ripples when using the traditional short-pulse excitation. The performance of this
technique was verified through simulation and experiments using the prototype integrated
HIFU/imaging transducer. Although it is known that the remnant ripples are related to the
135
notch attenuation value of the fixed notch filter, in reality, it is difficult to find the
optimal notch attenuation value due to the change in targets or the media resulted from
motion or different acoustic properties even during one sonication pulse. On the contrary,
the proposed adaptive noise canceling technique is capable of optimally minimizing both
the therapeutic interference and residual ripples without such constraints.
The prototype integrated HIFU/imaging transducer was composed of three
rectangular elements with cylindrically focusing. The 6 MHz center element was used for
imaging and the outer two identical 4 MHz elements worked together to transmit the
HIFU beam. Two HIFU elements of 14.4 mm × 20 mm dimensions with cylindrically
focusing could increase the temperature of the soft biological tissue from 55 ºC to 71 ºC
during 60 seconds. Two types of experiments for simultaneous therapy and imaging were
conducted to acquire a single scan-line and B-mode image with an aluminum plate and a
slice of porcine muscle, respectively. The B-mode image was obtained using the single
element imaging system during HIFU beam transmission. The experimental results
proved that the combination of the traditional short-pulse excitation and the adaptive
noise canceling method could significantly reduce therapeutic interference and remnant
ripples and thus may be a better way to implement real-time simultaneous therapy and
imaging.
In noninvasive HIFU treatment, formation of a large tissue lesion per sonication
is desirable for reducing the overall treatment time. This thesis showed the feasibility of
enlarging tissue lesion size with the DFTUT by increasing the DOF. The proposed
136
transducer is composed of a disc- and an annular-type element of different radii of
curvatures to produce two focal zones. To maintain uniform beamwidth of the elongated
DOF, inner and outer elements transmit ultrasound of a different center frequency, i.e.,
the inner element with a higher frequency for near field focusing and the outer element
with a lower frequency for far field focusing. The overall DOF of the proposed
transducer may be considerably extended by activating two elements simultaneously with
a single transmitter capable of generating a dual-frequency mixed signal.
A prototype transducer composed of a 4.1 MHz inner element and a 2.7 MHz
outer element was fabricated aimed at the treatment of malignant prostate tissues. The
performance of the prototype transducer was demonstrated through hydrophone
measurement, bio-heat transfer simulation, and tissue ablation experiment with a beef
liver. When several factors affecting the length and the uniformity of elongated DOF of
the DFTUT are optimized, the proposed therapeutic ultrasound transducer design may
increase the size of ablated tissues in the axial direction and thus decreasing the treatment
time for a large volume of malignant tissues. This technique can be applied to an
integrated HIFU/imaging transducer to accomplish simultaneous therapy and imaging
with reduced treatment time.
137
5.2 Further Works
Further work will focus on the fabrication of the IMCPA for treatment of the
malignant prostate tissue as shown in Figure 5.1. In this thesis, the prototype integrated
HIFU/imaging transducer used for demonstration of the proposed approaches was
composed of three single elements with the same dimension of the IMCPA with 64
elements for each array. To fully demonstrate the proposed techniques, an array type
transducer should be built and 64 elements for each array may be desirable based on the
preliminary experimental results. First, the flexible circuits for HIFU and imaging array
transducers need to be designed considering routing and bonding process based on RF
shielding and high temperature operation. Second, scratch dicing on the PZT4 and PZT-
5H with 1-3 composite structure may be carried out to separate each element before
bonding flexible circuits to each element. The specially designed housing with two
separated walls between HIFU and imaging transducer will be fabricated considering the
endocavity transducer. The IMCPA will be placed inside a cooling system using a water
tube to protect transducers from high temperature.
A 64-channel array system for imaging is preferred to monitor a hyperechoic
region in real-time using the IMCPA. The imaging system will be able to transmit both
coded and short pulse signals. The fixed and adaptive notch filtering function will be
implemented by a dedicated DSP for faster data processing time. Also, a 64-channel
138
HIFU array system will realize electronic scanning and dynamic focusing for the HIFU
array transducer in the IMCPA.
Figure 5.1: Schematic diagram for the integrated multi-functional confocal phased array
(IMCPA).
139
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Abstract (if available)
Abstract
Recently, high intensity focused ultrasound (HIFU) was successfully used for noninvasive treatment of the benign or malignant tissues. In ultrasound image-guided HIFU (US-gHIFU) using an integrated HIFU/imaging transducer, real-time simultaneous therapy and imaging is more desirable because it allows for tracking tissue movement and monitoring feedback induced by a treated target. However, reflected HIFU signals corrupt the quality of signals received by an imaging transducer during treatment.
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Jeong, Jong Seob
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Core Title
Transducers and signal processing techniques for simultaneous ultrasonic imaging and therapy
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Viterbi School of Engineering
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Doctor of Philosophy
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Biomedical Engineering
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04/27/2010
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