Design, development and characterization of wideband ... data illustrating the variation of...

Design, Development and Characterization of

Wideband Polymer Ultrasonic Probes

for Medical Ultrasound Applications

A Thesis

Submitted to the Faculty

of

Drexel University

by

Vadivel Devaraju

in partial fulfillment of the

requirements for the degree

of

Doctor of Philosophy January 2003

ii

DEDICATIONS

Dedicated to

MY FATHER

for his gentle humanity and dedication to

excellence in bringing up me to my present level

by taking the sufferings all along for my well-being.

In memory of MY MOTHER

for all her sacrifices in life for my betterment

and for her magnificent devotion to our family.

MY WIFE

a part of my soul – for her love, support, dedication

and commitment to our family value, culture and tradition.

MY SON

who is my strength, reflection of my thoughts and my emulation –

for his great affection, understanding, forbearance and

standing shoulder to shoulder with me in preserving our family heritage.

They are all the unbreakable four pillars of my life upon which I always stand unbowed.

ACKNOWLEDGEMENTS

iii

I express my sincere thanks and gratitude to my thesis advisor Dr.Peter A. Lewin for his

advice and support. My decade of association with Dr. Lewin will be evergreen in my

memory. My sincere thanks and appreciation from the bottom of my heart are to my

thesis committee members Dr. Richard B. Beard, Dr. T. S. Venkataraman, Dr. Hector V.

Ortega and Dr. Feroze B. Mohamed for their guidance and encouragement during the

process of my dissertation.

I also extend my thanks to Dr. Ryszard M. Lec for providing me the office space in his

lab and Dr. Philip E. Bloomfield for his assistance. I gratefully acknowledge the

financial support provided by Dr. Banu Onaral and Dr. William Freedman in the form of

Calhoun fellowship.

A debt of gratitude is owed to Dr. Mark E. Schafer for his support and for the help in

developing the new type of hydrophone and including it as a part of this thesis. I wish to

thank Dr.Hendrik Bleeker for all his help particularly while developing the low frequency

calibration method.

I express my gratitude to my brother, sisters and the members of their family for taking

my family burden back at home, and giving me the strength to face the challenges in life.

I express my thanks to all my in-laws for their timely help from time to time and to all

my relatives and friends who are all along with me by extending helping hand at all times

particularly during the time of need.

Table of Contents

iv

LIST OF TABLES………...……………………………………………………………...xi

LIST OF FIGURES………..…………………………………………………………….xii

ABSTRACT…………………………………………………………………………….xvi

CHAPTER 1. INTRODUCTION……………………………………………….………...1

1.1 Objectives, Outline and Scope of the Research…………………………….…2

1.1.1 Objectives…………………………………………………………...2

1.1.2 Outline ………………………………………………………………5

1.1.3 Scope of the research………………………………………………..6

1.2 Motivation……………………………………………………………………..7

1.3 Significance……………………………………………………………………8

1.4 Fundamentals of Biomedical Ultrasound…………………………………….10

1.5 Fundamentals of Medical Ultrasound Imaging………………………………12

1.6 Fundamentals of Medical Ultrasound Imaging Transducer………………….15

1.6.1 Resolution………………………………………………………….16

1.6.1.1 Axial resolution…………………………………………..16

1.6.1.2 Lateral resolution………………………………………...17

1.6.1.3 Transverse resolution…………………………………….19

1.6.1.4 Contrast resolution……………………………………….20

1.6.2 Bandwidth………………………………………………………….20

1.6.3 Sensitivity………………………………………………………….20

1.6.4 Backing material…………………………………………………...22

1.6.5 Electrical matching/tuning…………………………………………23

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1.6.6 Mechanical matching (Acoustic interface/impedance characteristics)……………………………………………………..23

1.6.7 Acoustic energy loss…………………………………………….…26

1.7 Fundamentals of Array Transducer………………………………………….27

1.8 Fundamentals of Ultrasonic Hydrophone Probe……………………………..29

CHAPTER 2. BACKGROUND AND THEORY……………………………………….31

2.1 Introduction…………………………………………………………………..31

2.2 Ultrasonic Polymer Hydrophone Probe……………………………………...32

2.2.1 Need for wider bandwidth hydrophone probe……..……………....34

2.3 Calibration Techniques………………………………………………………35

2.3.1 Planar scanning technique………………………………………….39

2.3.2 Time delay spectrometry technique combined with substitution method………………………………………………...39

2.3.3 Need for developing a new measurement procedure………………40

2.4 Single Element Imaging Transducer…………………………………………41

2.4.1 Transducer requirements…………………………………………...41

2.4.2 Conventional piezoceramic imaging transducers………………….42

2.4.3 Conventional polymer imaging transducers ………………………42

2.4.3.1 Single element, single layer transducer …………………42

2.4.3.2 Single element, multilayer polymer transducer …………43

2.4.4 Barker code polymer imaging transducer …………………………44

2.4.4.1 Barker code concept……………………………………...44

2.4.4.2 Principle of operation…………………………………….45

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2.5 Multielement (array) Imaging Transducer …………………………………..48

2.5.1 Side lobes and grating lobes……………………………………….49

2.5.2 Cross talk effects…………………………………………………...50

2.5.2.1 Electrical cross talk………………………………………51

2.5.2.2 Acoustical cross talk……………………………………..51

2.5.2.3 Electromechanical coupling……………………………...52

CHAPTER 3. PIEZOELECTRIC MATERIALS………………………………………..53

3.1 Introduction…………………………………………………………………..53

3.2 Material Properties…………………………………………………………...53

3.2.1 Piezoelectric stress constant………………………………………..54

3.2.2 Piezoelectric strain constant………………………………………..54

3.2.3 Transmitting constant……………………………………………....54

3.2.4 Receiving constant………………………………………………....55

3.2.5 Dielectric permittivity (constant)…………………………………..55

3.2.6 Electromechanical coupling coefficient……………………………55

3.2.7 Dissipation factor…………………………………………………..56

3.2.7.1 Mechanical loss tangent……………………………….…56

3.2.7.2 Electrical loss tangent……………………………………57

3.2.8 Acoustic impedance………………………………………………..57

3.3 Piezoelectric Materials for Imaging Transducers …………………………...57

3.3.1 Piezoceramic……………………………………………………….57

3.3.2 Piezocomposite…………………………………………………….59

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3.3.3 Single crystal………………………………………………………60

3.3.4 Piezopolymer………………………………………………………62

3.4 Modes of Vibration…………………………………………………………..65

3.5 Summary……………………………………………………………………..66

3.6 Conclusion…………………………………………………………………...68

CHAPTER 4. DEVELOPMENT AND CHARACTERIZATION OF IMPROVED DESIGN OF POLYMER HYDROPHONE PROBE ……………………69

4.1 Introduction…………………………………………………………………..69

4.2 Synopsis……………………………………………………………………...70

4.2.1 Acoustic sensitivity………………………………………………...70

4.2.2 Frequency response and bandwidth………………………………..70

4.2.3 Angular response…………………………………………………..71

4.2.4 Effective aperture size……………………………………………...71

4.2.5 Orientation effects………………………………………………….71

4.3 Conclusion…………………………………………………………………...72

CHAPTER 5. DEVELOPMENT OF CALIBRATION PROCEDURE……….………...73

5.1 Introduction…………………………………………………………………..73

5.2 Calibration Setup…………………………………………………………….73

5.3 Calibration Procedure………………………………………………………..75

5.3.1 Initial alignment procedure………………………………………...75

5.3.2 Determination of the peak pressure amplitude location …………..76

5.3.3 Calculation of pulse intensity integral …………………………….76

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5.3.4 Cross axis scan……………………………………………………..78

5.3.5 Raster scan…………………………………………………………79

5.4 Radiation Force Balance Measurements……………………………………..83 5.5 Hydrophone Sensitivity Calculation…………………………………………84

5.6 Comparative Study…………………………………………………………...85

5.7 Results Correlation and Discussion………………………………………….86

5.8 Summary……………………………………………………………………. 88

5.9 Conclusions…………………………………………………………………..89

CHAPTER 6. DEVELOPMENT AND CHARACTERIZATION OF SINGLE ELEMENT NON-RESONANT POLYMER IMAGING TRANSDUCER ………………………………………………………….91

6.1 Introduction…………………………………………………………………..91

6.2 Characterization of PVDF Film – Impedance / Admittance Measurements...91

6.3 Design and Development of System Electronics and Controls……………...92

6.4 Design Considerations of Single Element Polymer Imaging Transducer…...93

6.4.1 Material selection, thickness and active area…………………..93

6.4.2 Electroding……………………………………………………..94

6.4.3 Backing material……………………………………………….94

6.4.4 Wear protecting front matching layer………………………….95

6.4.5 Adhesion……………………………………………………….96

6.5 Fabrication Process…………………………………………………………..96

6.6 Characterization of Non-Resonant Single Element, Single Layer Transducer and Multilayer Barker Code Polymer Imaging Transducer…...101

6.6.1 Electrical characterization…………………………………….101

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6.6.2 Experimental system………………………………………….102

6.6.3 Pulse-echo response…………………………………………..104

6.6.4 Diffraction correction…………………………………………108

6.6.5 Sensitivity correction…………………………………………113

6.7 Summary……………………………………………………………………114

6.8 Conclusion………………………………………………………………….115

CHAPTER 7. DESIGN, DEVELOPMENT AND CHARACTERIZATION OF NON-RESONANT MULTI ELEMENT (ARRAY) SINGLE LAYER/MULTILAYER POLYMER IMAGING TRANSDUCER…...117

7.1 Introduction………………………………………………………………....117

7.2 Design Consideration and Description of Array Structure…………………117

7.3 Fabrication Process…………………………………………………………120

7.4 Design and Development of System Electronics and Control……………...124

7.5 Experimental System for Performance Evaluation…………………………124

7.6 Performance Evaluation of the Array Transducer …………………………126

7.6.1 Single layer array transducer………………………………….….126

7.6.1.1 Pulse echo response of individual element……..………126

7.6.1.2 Pulse echo response of all the four elements……..…….128

7.6.1.3 Uniformity………………………………………………129

7.6.2 Multilayer array transducer ……………………………….……...130

7.6.2.1 Pulse echo response…………………………………….130

7.7 Summary ……………...……………………………………………………132

7.8 Conclusion………………………………………………………………….133

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CHAPTER 8. SUMMARY AND CONCLUSION…………………………………….134

8.1 Summary of the Research Work ……………………...……………………134

8.2 Conclusion………………………………………………………………….136

8.3 Suggestion for Future Work………………………………………………...138

BIBLIOGRAPHY………………………………………………………………………140

APPENDIX A: DEVELOPMENT AND CHARACTERIZATION OF IMPROVED DESIGN OF DOUBLE LAYER POLYMER HYDROPHONE PROBE………………………………………………………………...147

VITA……………………………………………………………………………………183

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List of Tables

1. Comparison of piezoelectric material parameters…………………………….……….67 2. Comparative statement showing the variation of frequency response and sensitivity of hydrophones having different thickness of PVDF film………………167 3. Directivity data illustrating the variation of effective diameter of the hydrophones measured along the two orthogonal axes at 5, 7.5 and 10 MHz………174 4. Directivity data illustrating the variation of mean effective diameter of the hydrophones measured along the two orthogonal axes……………………………...174 5. Directivity data illustrating the variation of effective diameter of the hydrophones having different thickness of PVDF film measured along the two orthogonal axes……………………………………………………….178

xii

List of Figures

1. Schematic representation of the outline of the research work…………………………5 2. Schematic representation of the Barker code arrangement…………………………...45 3. Geometry of three important vibration modes……………………….……………….65 4. Typical measured hydrophone waveform with the corresponding PII and frequency spectrum of 1 MHz circular piston source……………………………77 5. Beam plot along the x and y axes of the 1 MHz acoustic source. The dashed lines show the gaussian beam (theoretical) distribution with the same –6 dB beamwidth as that of the measured transducer………………...79 6. Experimental set up for rater scanning (A) and illustration of raster fashion (B)……81 7. Three-dimensional representation of the intensity field produced by the 1 MHz source transducer and obtained from planar scanning…………………….….82 8. End-of-cable voltage sensitivity versus frequency of the double layer PVDF hydrophone in the frequency range 0.3 to 1 MHz obtained using Planar scanning technique (), Time delay spectrometry method (—) and calibration data provided by National Physical Laboratory (NPL, ◊), UK……….….86 9. Plot showing the variation of admittance magnitude and phase of a 56 µm piezo film in air………………………………………………………………………92 10. Plot showing the variation of impedance magnitude and phase of a 56 µm piezo film in air……………………………………………………………….……...92 11. Pressing mechanism while in use……………………………………………………98 12. Different stages of the fabrication process of single element transducer…………..100 13. Completed single element transducers……………………………………………...100 14. Plot showing the variation of admittance magnitude and phase of single layer transducer ……………………………………………………………………101 15. Plot showing the variation of impedance magnitude and phase of a single layer transducer …………………………………………………………………….101

xiii

16. Plot showing the admittance magnitude of 56 µm piezo film, one layer and three layer transducer ………………………………………………………….101 17. Experimental setup for pulse echo measurement…………………………………...102 18. Pulse-echo response (in time and frequency domain) of single element, single layer transducer at 1 cm depth in water for a monocycle sine burst excitation……………………………………………………………………………104 19. Experimental pulse echo response (in time domain) of a single element, three layer Barker code transducer at 1 cm depth in water for a monocycle sine burst excitation………………………………………………………………...105 20. Pulse echo frequency response of single element, single layer transducer at different depths in water for a monocycle sine burst excitation ………………...105 21. Pulse-echo frequency response of single element, three-layer transducer at different depths in water for a monocycle sine burst excitation…………………105 22. Pulse-echo frequency response of single and three-layer PVDF transducers at 1 cm depth in water for a monocycle sine burst excitation………………………106 23. Pulse-echo frequency response of one single layer and two three layer single element polymer transducers at 1 cm depth in water for a monocycle sine burst excitation………………………………………………………………...108 24. Pulse echo response (in time and frequency domain) of single element, single layer polymer transducer at different depths in water for monocycle sine burst excitation………………………………………………………………...111 25. Pulse echo response of a single element, single layer transducer for different water path length at peak frequency for monocycle sine burst excitation………….112 26. Pulse echo response of single layer transducer at different depths in water for a monocycle sine burst excitation (with and without diffraction correction)………...112 27. Schematic diagram showing the capacitance loading of the polymer transducer …113 28. Schematic representation of array element and bonding pad pattern at one side of the polymer film of the proposed linear array transducer ………………….119 29. Actual electrode pattern of an array layer…………………………………………..121 30. Different stages of the construction process of 4 element array transducer………..123 31. Completed array transducers………………………………………………………..123

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32. Photograph of the newly built transmit/receive control circuit and preamplifier…..124 33. Experimental configuration for pulse echo measurements……….………………...125 34. Experimental pulse echo responses (in time & frequency domain) for one of the elements of single layer array transducer at 1 cm depth in water for a monocycle sine wave excitation……………………………………………………127 35. Experimental pulse echo response (in time and frequency domain) while exciting all the four elements of single layer array transducer at 1 cm depth in water for a mono cycle sine wave excitation……………………………...128 36. A typical observed time domain pulse echo response of one of the elements of the single layer array transducer at 1 cm depth in water for a monocycle sine burst excitation………………………………………………………………...129 37. Pulse-echo responses (in time & frequency domain) of a representative element stack of three layer array transducer at 1cm depth in water for a monocycle sine wave excitation………………………………………………………………...130 38. Typical observed time domain pulse echo response of one of the elements of single layer (A) and one of the element stacks of three layer (B) array transducer at the peak frequency at 1 cm depth in water for a monocycle sine burst excitation………………………………………………………………...131 39. Pulse echo frequency response of two element stacks of three layer array transducer at 1 cm depth in water for a monocycle sine burst excitation…………..132 40. Actual view of the newly developed double layer polymer hydrophone probe……160 41. Schematic representation showing the capacitance loading of the hydrophone probe………………………………………………………………………………..164 42. Frequency response plot of a double layer hydrophone probe using 9+9 µm thick polymer film…………………………………………………………………..169 43. Frequency response plots of double layer hydrophones using 25+25 µm and 25+9 µm polymer film……………………………………………………………...170 44. Typical directional response of double layer polymer membrane hydrophones with active element size of 0.4 mm (Fig. a), 0.6 mm (Fig. b) and 1 mm (Fig. c) in diameter, measured at the frequencies of 5 MHz, 7.5 MHz and 10 MHz………………………………………………………………175

xv

45. Combined directional response of double layer polymer membrane hydrophones with active element size of 0.4 mm in diameter, measured at the frequencies of 5 MHz, 7.5 MHz and 10 MHz…………………….177 46. Combined directional response of double layer polymer membrane hydrophones with active element size of 0.4 mm, 0.6 mm and 1 mm in diameter, measured at 7.5 MHz…………………………………………...177 47. Combined directional response of double layer polymer membrane hydrophone with active element size of 0.5 mm in diameter, measured at 10 MHz in two orthogonal axes………………………………………………….178

xvi

Abstract Design, Development and Characterization of

Wideband Polymer Ultrasonic Probes for Medical Ultrasound Applications

Vadivel Devaraju Peter A. Lewin, Ph.D.

This dissertation deals with the design, development and characterization of non-resonant

polymer ultrasonic probes for medical ultrasound applications. Both single element and

multielement imaging transducer design having single layer and multilayer configuration

were developed with the primary goal of minimizing the trade off between resolution and

penetration depth.

The simultaneous improvement in the transducers’ pulse-echo sensitivity and bandwidth

was achieved by employing a multilayer structure made of thin piezopolymer films and

utilizing the concept of Barker code. The results of the experiments indicated that the

multilayer Barker code transducers provided the widest bandwidth in comparison with

the imaging transducers made of conventional piezoelectric ceramic material. Also, they

exhibited enhanced sensitivity compared to a single layer piezopolymer transducer.

Specifically, the results indicated that the –6 dB fractional bandwidth extending over 2

decades (20 MHz) could be achieved in the case of non-resonant transducers, whereas the

conventional resonant design imaging transducer could provide only about one half of the

bandwidth. The polymer array transducers also showed uniform acoustic response from

element to element, which is desirable in order to obtain high quality ultrasound images.

xvii

The double layer hydrophone probes used for characterization of the imaging transducers

were fabricated using dissimilar thickness of PVDF polymer film. This technique

ensured simultaneous enhancement of sensitivity and bandwidth. Also, a measurement

procedure employing planar scanning technique was developed to calibrate the

hydrophone probes below 1 MHz for adequate characterization of acoustic field produced

by the imaging transducers.

The results of this work indicate that the new class of transducers developed features

significantly enhanced bandwidth. Such transducers hold promise to be capable of

operating at clinically relevant frequencies and suitable for use at fundamental,

subharmonic and higher harmonics imaging. It is expected that this non-resonantly

operating imaging transducer would become a useful clinical tool in medical imaging and

could improve diagnostic efficacy.

1

CHAPTER 1: INTRODUCTION

This dissertation deals with the design, development and characterization of wideband

polymer ultrasonic probes for medical ultrasound applications. The research work

included both the single element and multielement (array) imaging transducers design

and also hydrophone probe used in characterization of the imaging transducers. The first

phase of this work involved fabrication of a double layer hydrophone using dissimilar

thicknesses of polyvinylidene fluoride (PVDF) polymer film and development of

associated measurement procedures. Next, the experience and knowledge gained in the

fabrication process were used to design and develop the single element, multilayer

imaging transducers and multielement (array), multilayer imaging transducers. The

overall goal of this work was to investigate a possibility to minimize the unavoidable

trade off between the resolution and penetration depth in clinical ultrasound imaging

applications.

The thesis is divided into eight chapters. This chapter reviews the objectives and the

scope of the work performed, and the significance and motivation for the research along

with a brief review of fundamentals of the ultrasound essential for understanding of the

tasks involved in development of the non-resonant imaging probes. Chapter two reviews

background and theory needed in the course of this work and also provides the

background information including review of the work done by other researchers. Chapter

three contains detailed information about different properties of various piezoelectric

materials used in imaging probes and justifies the effort in exploring the applicability of

2

the polymer as piezoelectric material for imaging transducers. A brief review of the

development and characterization of double layer PVDF polymer hydrophone probes

used for characterization of the imaging transducers developed in this work is given in

chapter four, whereas a more detailed description of the development process and results

are presented in Appendix A. Chapter five deals with the measurement procedure

developed for calibrating the double layer ultrasound hydrophone probe below 1 MHz

using planar scanning technique along with the related background and the calibration

results. Chapter six describes the development of a single element, single layer

transducer and three layer Barker code transducers. The design and development of

multielement, single layer as well as multilayer array transducers are presented in chapter

seven. Finally, chapter eight presents the summary of the work accomplished together

with conclusions and suggestions for future work.

1.1. Objectives, Outline and Scope of the Research

1.1.1 Objectives

In order to achieve the primary goal of having a single imaging ultrasound transducer

capable of operating at all clinically relevant frequencies and minimizing the trade off

between image resolution and penetration depth, a set of objectives was devised. The

objectives of the research described in the following were to design, develop and

characterize piezoelectric polymer ultrasonic probes used for medical ultrasound

applications, which are capable of exhibiting simultaneous enhancement of both the

bandwidth and sensitivity in the case of both imaging transducers and hydrophone

probes. The objectives were accomplished by performing the following specific tasks:

3

Specific Task (1): Development of non-resonant single element, single layer transducer and multilayer Barker code imaging transducers

In this phase the focus was on the feasibility of implementation of a non-resonant

transducer with maximized bandwidth using Barker code and verification of this concept

based on the experimental results. The design used 56 µm thick PVDF film, resulting in

the development of a new class of non-resonantly operating wideband polymer

transducers that might be usable at clinically relevant frequencies (2-15 MHz) using a

single imaging transducer. The performance of the transducers was verified by

characterizing them in pulse-echo mode.

Specific Task (2): Design, development and characterization of non-resonant multielement single layer and multilayer linear array imaging transducers

A feasibility study was performed to demonstrate the Barker code concept with simple,

multielement single layer and three-layer linear array transducers by extending the work

mentioned in specific task (1). The linear array transducer was developed using 56 µm

thick PVDF film having four elements and its characterization included (a) measurement

of the pulse-echo response of individual element (b) test of the uniformity of the response

and (c) recording of the spectral response of echoes from each element and element stack

of the array.

4

Specific Task (3): Fabrication and characterization of wideband double-layer pvdf Polymer hydrophone probes and development of measurement procedure

In the initial phase of this research, double layer piezoelectric polymer hydrophone

probes used for acoustic measurements of imaging transducers were fabricated using

identical and different thicknesses of PVDF film. Their performance characteristics were

investigated by experimentally determining the frequency response, directional response,

and effective aperture size at two different orientations (planes).

A measurement procedure to calibrate the hydrophone probe below 1 MHz for adequate

characterization of the acoustic field produced by the imaging transducers was also

developed. The procedure was based on Planar Scanning Technique [1] and the results

of the measurements were verified with the data obtained by using Time Delay

Spectrometry (TDS) method [2, 3, 4] and also with the data provided by National

Physical Laboratory (NPL), United Kingdom [5].

5

1.1.2 Outline

The outline of the research work is schematically represented in Figure 1 and a detailed

explanation is presented in the next section namely, the scope of the research.

Phase 1

Phase 2

Phase 3

Figure 1: Schematic representation of the outline of the research work

Ultimate Goal

Single Imaging Probe for all diagnostic purposes

Design & Development of Multielement (Array)Single/MultilayerImaging Transducer

Development of Single ElementSingle/MultilayerImaging Transducer

Fabrication andCharacterization ofDouble Layer PVDFHydrophone Probe

Development ofMeasurememtProcedure

Bottom Up

Specific Task

Specific Task

Specific Task

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6

1.1.3 Scope of the research

As shown in Figure 1, the research work presented here involved: (1) development of a

wideband single element single/multilayer imaging probe, (2) design and development of

a wideband multielement (array) single/multilayer probe as a feasibility study, (3)

development of double layer hydrophone probe and (4) development of measurement

technique and procedure. All four tasks of this research work are closely interrelated.

The scope of the work also involved experimental evaluation of newly developed

ultrasound transducers and critical review and discussion of the various results.

The primary contribution of this work is (a) development and testing of nonresonant

multilayer imaging transducers in pulse echo mode. Specifically, the possibility of using

the Barker code technique to enhance desirable properties of imaging transducers is

demonstrated; (b) design and development of multielement (array) multilayer transducer

to validate the Barker code concept. (c) development of a double layer ultrasound

hydrophone probe using dissimilar thickness films that resulted in simultaneous

enhancement of its bandwidth and sensitivity; and (d) optimization of the planar scanning

technique for calibration of the hydrophones below 1 MHz.

A transmit/receive control system and preamplifier were designed and fabricated to

evaluate the imaging probes. A pressing mechanism [6] was constructed and the

thickness of the bond line was optimized in laminating the multilayer structure. In

summary, this thesis presents a possible solution to the problems the clinicians are facing

at present. These include: the difficulty in using a single imaging transducer for all

7

diagnostic applications and the trade off between the image resolution and penetration

depth.

1.2 Motivation

As noted above, the primary motivation of this research was prompted by a need to have

an imaging transducer, which would provide both adequate pulse echo sensitivity and

broad bandwidth, concurrently. This would allow a single scanhead to be used in clinical

diagnostic practice. The existence of such scanhead would shorten the examination time,

allow immediate, on-the-site optimization of image resolution, and hence improvement in

diagnostic efficacy.

By studying the properties of the several types of piezoelectric material, it was decided to

explore the feasibility of using piezopolymer for the imaging transducer because of the

reasons outlined in the following. Although conventional piezoceramic material provides

good sensitivity, it suffers from a narrow bandwidth. The PZN/PT single crystal

demonstrated good pulse-echo sensitivity and relatively wide bandwidth in comparison to

solid (PZT) ceramic and piezocomposite [7]. However, the widest possible bandwidth

could be realized by using a piezopolymer such as PVDF. Although this piezopolymer

exhibits inadequate pulse echo sensitivity, by employing the multilayer structure and

utilizing the concept of Barker code, acceptable sensitivity and widest possible bandwidth

could be simultaneously achieved. The literature search revealed that no one has

developed the Barker code array transducer for medical ultrasound applications so far.

8

1.3 Significance

In medical ultrasound imaging, the overall system performance is mainly limited by the

imaging transducer characteristics. Consequently, there is a strong desire to further

improve the sensitivity and bandwidth of such transducers. Advanced signal processing

techniques, which use frequency modulated functions rely upon the transducers’

possessing good sensitivity and bandwidth characteristics. It is well known that the axial

resolution (i.e., the ability to distinguish closely related reflecting structures, which lie in

different planes parallel to the probe face) of an image is related to the center frequency

and bandwidth of the probe. High lateral resolution (i.e., the ability to distinguish closely

related reflecting structures, which lie in different planes perpendicular to the transducer

face) requires also wide frequency bandwidth besides other design parameters. The

bandwidth of conventional imaging probes made of solid or composite ceramic and used

in diagnostic systems is limited due to the fact that they operate at the resonance.

Therefore, to optimize the image resolution and penetration depth, multiple transducers

operating at different resonance frequencies are being used at present to cover the

frequency range of interest during diagnosis due to the relatively narrow bandwidth of the

PZT (Lead Zirconate Titanate) ceramic transducers. The PZT transducers consist of a

thin disk, operated at resonance and the bandwidth is determined by the thickness

resonance frequency (λ/2) of the ceramic. Because of the brittleness of the PZT material,

thin disks are difficult to use during the imaging transducer fabrication procedure. As the

PVDF is available as thin film, the polymer transducers can be operated at frequencies

well below the resonance (thickness resonant frequency of 9 µm thick film is about 122

MHz).

9

The majority of medical ultrasound systems utilize real time imaging arrays and arrays

operating at frequencies up to about 15 MHz in clinical practice. Beyond 15 MHz, there

is a practical limitation in manufacturing of multi-element arrays using solid PZT

material, as the required element dimension is difficult to achieve in the fragile piezo-

ceramic materials. Besides, there is the associated problem with the necessity of using

matching layers. Hence, there is a need to develop new classes of wideband ultrasonic

transducers for medical imaging capable of operating at clinically relevant frequencies

and providing simultaneous improvement of both the bandwidth and sensitivity.

It is conceivable that instead of using several ceramic transducers necessary to cover the

frequency range of different clinical ultrasound applications, if properly designed, a

single PVDF probe with wider bandwidth could be used. Despite this advantage, with a

few exceptions, PVDF material is not used in the design of imaging transducers due to its

relatively low electromechanical coupling coefficient and dielectric constant, resulting in

a poor transmitting efficiency. The exceptions include a few applications in the field of

skin and ophthalmology, wherein single element PVDF probes are employed. So far,

multilayer non-resonant PVDF array transducers have not been used in medical imaging.

Accordingly, the research described here focused on a feasibility study, which involved

development of a multilayer linear array polymer imaging transducer incorporating the

concept of Barker code. The study resulted in fabrication and testing of a new class of

imaging transducers featuring enhanced bandwidth and acceptable pulse-echo

performance.

10

As already noted, the transducer designed once optimized, will be potentially capable of

operating at any clinically relevant frequency. So, one transducer could be sufficient for

most of the diagnostic imaging. The existence of such a wideband transducer would

shorten patient’s examination time, allow immediate, on-the-site optimization of image

resolution and lead to improvement in diagnostic efficacy. In the next section

fundamentals of biomedical ultrasound, medical ultrasound imaging and imaging

transducers are outlined. This is done to facilitate a better understanding of the work

presented.

1.4 Fundamentals of Biomedical Ultrasound

Sound is the experience of the propagation of pressure through some physical elastic

medium such as liquid or air. The pressure waves are generated due to some type of

mechanical disturbance. The vibrating pressure waves transfer the mechanical energy in

the form of wave to the medium and to objects that the wave contacts. If the vibrational

frequency is very high, human cannot hear it, as the human hearing is limited up to about

20 kHz. There are mammals such as bats and whales, which can sense above 100 kHz to

determine distance. The ultrasound is a mechanical wave with a frequency above the

limit of human hearing. In general, there is a fundamental higher limit of frequency

based on the closest atomic spacing in solids, which is about 1013 Hz.

Using ultrasound, the interior of an object can be examined by transmitting pulses of high

frequency sound waves into the object and receiving the echoes from the internal

structure by the same source transducer. The time delay for the reflected echo to return

11

to the source will determine the distance of the reflecting structure. The pulse echo

technique was originally developed for radar and sonar. In 1917 Langevin in France used

the pulse echo technique to measure the depth of water in the ocean. After World War II,

J.M.Reed and J.J.Wild first adopted the pulse echo techniques to image biological

structure in 1952 [8].

An ultrasound transducer made with piezoelectric material translates the applied voltage

pulses into high frequency sound vibration, and then listens for the echoes. The

transducer converts the echoes (pressure waves) into voltage and the waveform is

converted into digital representation, which is then converted into an image by electronic

devices so as to be seen on a monitor. Medical ultrasound is a process, where the sound

waves are bounced into the body and the reflected sound waves are captured and

transformed into an image by the ultrasound machine.

The resolution of the image is fundamentally limited by the wavelength of the ultrasound.

The resolution increases with increasing frequency. As the ultrasound travels through the

biological material, it gets attenuated due to spreading, scattering and absorption. The

rate of attenuation in tissue is proportional to the frequency of ultrasound. Therefore,

there is a trade off between resolution and penetration depth. The frequency of

ultrasound needs to be high enough to give good resolution and low enough to detect the

echoes from the deeper structures. So the constraints on frequency control the scale of

the anatomy to be examined. Currently, frequencies from 1 to 15 MHz are generally

used in medical diagnosis except in some special cases such as in the field of

12

ophthalmology and dermatology where higher frequency single element transducers are

used. Transducers operating at 250 kHz to 1 MHz are commonly used in many

therapeutic applications.

1.5 Fundamentals of Medical Ultrasound Imaging

In a diagnostic ultrasound system, the ultrasound pulses are generated by the imaging

transducer containing a piezoelectric material. The transducer is used both for generating

the pulses and detecting the echoes. When the piezoelectric material is subjected to an

electrical voltage, it undergoes a change in dimension. If the electrical stimuli are of very

short duration, the piezoelectric material experiences very rapid fluctuations in

dimension. These rapid mechanical vibrations produce ultrasonic pressure waves.

Conversely, if the pressure wave strikes the piezoelectric material, it causes mechanical

deformations of the piezoelectric material, which produce electrical voltage.

The ultrasound beam consists of train of pulses that are emitted by the probe about a

thousandth of a second apart (PRF ~ 1 kHz). Each pulse travels in the beam at the speed

of sound and is reflected by the structure within the body. The reflected sound beam

(echoes) gives information about the position of the reflecting structure. The pulse-echo

information from the reflecting structure can be displayed in different ways. One of the

ways is to find the distance, which is called A-Scan. In an A-scan display, the horizontal

axis represents the time required for the echo to return to the imaging transducer, which

corresponds to the distance between the imaging transducer and the reflecting organ. The

vertical axis shows the height of the peak amplitude, which relates to the strength of the

13

reflected echo. When ultrasound pulses travel through the human body, they are partially

reflected at boundaries between structures that differ in their characteristic impedance.

Since the difference between the impedance of one tissue with the other is small, the

echoes reflected by the boundaries of two tissues are faint. The boundaries between

tissue and bone return strong echoes due to large difference in acoustic impedance. The

echoes from the far side of an organ are fainter than the ones from the near side, due to

attenuation of ultrasound as it propagates deeper into the body. The frequency of the

reflected ultrasound wave from a stationary structure is equal to the frequency of the

incident wave.

The echo ranging information of the A scan can be displayed in the brightness modulated

form called B-Scan. In a B-Scan, echo is represented by a spot of light. The brightness

of the spot of light represents the strength of echo received. The position of the spot of

light represents the time required for the echo to return to the transducer and shown on a

time scale. The position and the brightness of the echoes are stored in the ultrasound

machine and the integrated image of all the stored signals forms a representation of

structure/organ in cross section.

In the case of moving structures such as blood or a beating heart, the reflected wave has a

different frequency than the incident wave. The shift in frequency is due to the Doppler

effect. The movement of the structure towards the imaging transducer compresses the

wavelength of the reflecting wave, causing an increase in the frequency. The movement

of the target away from the imaging transducer lengthens the wavelength of the reflected

14

wave, resulting in a decrease in the frequency. The shift in the frequency is proportional

to the frequency of the incident wave and the velocity of the moving structure. For

Doppler shift measurements, the ultrasound is transmitted in the form of a continuous

wave or quasi-continuous wave. The imaging transducer consists of two transducers, one

for emitting continuous ultrasound wave towards the moving object and another for

detecting the echoes from the target and also from the stationary structures surrounding

the intended target. The received signal is electronically mixed with the reference

frequency of the transmitting transducer. By mixing the two sound waves, four

frequency components are obtained namely frequency of the transmitted wave, reflected

wave, sum and difference of the two frequencies. The difference of the two frequencies

alone is filtered out which is called Doppler shift frequency. If the imaging transducer is

excited while aiming towards a stationary structure, no sound will be heard. When the

transducer is aimed at a moving object, the diagnostician can hear a tone/sound whose

frequency is equal to the Doppler shift frequency. The blood flow can be detected from

the reflected or backscattered ultrasound from the moving blood cells. For the beating

heart, the sound heard will vary with heartbeats. Since the Doppler system relies on

continuous wave propagation, it cannot measure the distance of the target from the

imaging transducer. By adopting a technique called gated transmission by time coding

and using a long pulse, distance information is obtained.

The state of the art ultrasound system uses imaging arrays instead of single element

transducers. The array consists of several transducing elements which are electronically

switched, steered and focused both in transmit and receive mode to form the high-

15

resolution images. A group of elements typically four or eight are generally used in

transmitting the ultrasound. The size of the group of elements excited determines the size

of the aperture and the shape of the transmitted sound beam. Each element of the array is

excited with a time delay signal in order to achieve a focused beam. The beam former in

the ultrasound machine selects the aperture, applies time delay and steers the beam. After

the echo signal leaves the receive beam former, it undergoes time gain compensation

(TGC), which compensates for the loss in amplitude of the signal due to depth and also

for the losses due to diffraction of ultrasound as well as the attenuation in tissue.

Logarithmic compression is employed to reduce the dynamic range of the signal so that

both the weak and strong echoes are within the range of 40 dB in order to display the

image on a monitor, which has a limited dynamic range. The next stage is the envelope

detection, which demodulates the echo signal and filters out the carrier signal. The

demodulated signal is sent to the scan converter. The above process illustrates one scan

line and the image display requires between 48 and 196 scan lines and a frame rate of

about 30 per second. Besides A-scan, B-scan and Doppler imaging, the recent advances

include tissue/native harmonic imaging, contrast imaging, 3 and 4 dimensional imaging,

coded excitation and elastography.

1.6 Fundamentals of Medical Ultrasound Imaging Transducer

Imaging transducers are constructed from a piezoelectric material so that the element will

resonate mechanically in a thickness mode at its fundamental frequency of oscillation.

The thickness is chosen to be half wavelength. An ideal imaging transducer should

achieve highest sensitivity and penetration, optimum focal characteristics and best

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possible resolution, all at low power output, conforming to ALARA (As Low As

Reasonably Achievable) principle.

1.6.1 Resolution

The resolution in medical imaging consists of four components namely (a) Axial

resolution (b) Lateral resolution (c) Transverse resolution (d) Contrast resolution.

1.6.1.1 Axial resolution

The axial resolution is defined as the minimum distance between two reflecting

structures, which lie in different planes parallel to the transducer face, at which they may

be distinguished as separate targets. The factors that contribute the axial resolution of an

ultrasound imaging systems are (i) nature of the pulsing and receiving circuit (ii) signal

processing (iii) transducer characteristics. The transducer parameters, which influence

the axial resolution, are (a) damping (ring down) (b) frequency (wavelength) / spatial

pulse length. A brief explanation about these parameters, which are pertinent to this

transducer development project, is outlined below for the sake of understanding of the

readers.

Damping

The vibration of the piezoelectric material should be damped out as quickly as possible.

As long as the material is vibrating or ringing, it is not available to receive the ultrasound

echoes from various reflecting structures. There is a trade off between limiting the

amount of ringing to improve the axial resolution and allowing it ring long enough to

provide sufficient acoustic energy in order to obtain echoes of sufficient magnitude. The

17

degree of damping or ring down gives a measure of the axial resolution potential. Highly

damped transducer will have lower sensitivity.

Frequency

The frequency of the transducer determines the axial resolution potential as well as the

penetration depth and lateral resolution. Increase in frequency improves the axial

resolution. It also gives better lateral resolution in view of less divergence of the beam.

Increase in frequency has decreased wavelength, which results in decreased spatial pulse

length.

Spatial pulse length

The spatial pulse length represents the actual physical space occupied by the burst of

ultrasound. It is measured from the start of the ultrasound burst to the –20 dB ring down

point. By the time the vibrations have been damped down to –20 dB, if the wavefront

has already traveled say, 6 mm, then the spatial pulse length is said to be 6 mm. So this

first 6 mm of the tissue cannot be visualized and also the reflecting structures separated

by less than 6 mm cannot be resolved.

The relationship is stated as follows: increased frequency decreased wavelength

decreased spatial pulse length better axial resolution.

1.6.1.2 Lateral resolution

If the ultrasound beam is narrow enough to pass between two reflecting structures

without being incident upon either one, then the beam is capable of resolving them as two

distinct structures. The lateral resolution is defined as the minimum distance between

two reflecting structures on a line perpendicular to the imaging beam at which they are

18

distinctly visualized as separate structures. The beamwidth of the ultrasound beam is the

parameter, which determines the lateral resolving capability of the imaging transducer.

As beamwidth increases, lateral resolution and image details deteriorate. Since the

ultrasound beams from single element transducers are circular at any one point, the

beamwidth is actually the beam diameter. It is determined by measuring the width of the

amplitude at –6 dB level. The transducer parameters that affect the beamwidth are (i)

geometry of the piezoelectric material (ii) frequency (iii) focusing (iv) distance from the

probe. The effect of these parameters on beam width and the consequent influence on

image quality are delineated below.

Geometry of the piezoelectric material

Circular elements produce cylindrical beams and hence the lateral resolution capability is

uniform across any cross section of the beam. The rectangular elements as used in linear

arrays produce elliptical beams and hence the lateral resolution is better in one plane of

the beam than the other. The element diameter affects the width and shape of the beam.

As element diameter increases, beam spread decreases, resulting in very narrow beams.

The far field divergence is given by the expression sinθ = 1.22λ/d, where d is the

diameter of the transducer. Small element diameter which produces narrow beam is

useful in examining superficial structures because reduced diameter will have reduced

beam diameter in near field and increased far field divergence, which has no consequence

in superficial exam. Larger element diameter will have reduced far field divergence,

resulting in improvement in imaging deeper structures.

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Frequency

Frequency also affects the beamwidth. High frequency beams are less divergent.

Increase in frequency causes an increase in the near field distance and decrease in far

field divergence. The combined effect produces an effectively narrower beam.

Focusing

Focusing is a means of regulating the width of the sound beam and the region of

maximum sensitivity. Focusing helps to improve the image, because it creates a

narrower beam over a defined focal zone. The focal point can only be adjusted within the

near field, while the focal zone may extend into the far field. The radius of curvature, in

combination with frequency and active element diameter, determines the focal point of

the transducer. The focal zones are defined as the region of sound beam where the

acoustic sensitivity falls to ½ (-6 dB) of the axial maximum. The selection of focal zone

for an imaging transducer is based on the depth of the tissue of interest so as to put the

narrowest portion of the beam in the region under study.

The active element diameter, frequency and focal zone must be taken together while

selecting an imaging transducer for a specific application, as each one directly affects the

beamwidth and consequently the lateral resolution and ultimately the image quality.

1.6.1.3 Transverse resolution

The transverse resolution is the directionality of the ultrasound beam in the transverse

direction and it depends on the transducer’s aperture, focusing and axial resolution.

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1.6.1.4 Contrast resolution

The contrast resolution is the minimum scattering level at which adjacent region can be

visualized as having a varied intensity level.

1.6.2 Bandwidth

In an ultrasound beam, the distance between two points at which the acoustic pressure

drops to 6 dB of the maximum value is defined as the bandwidth. A broad bandwidth

implies a shorter spatial pulse length, assuring excellent axial resolution. A broader

bandwidth results in larger area under the frequency distribution curve and therefore

more energy in the pulse, which translates to more sensitivity. The broad bandwidth

probe can be more precisely matched to the electrical characteristics of the ultrasound

(system) machine with which it operates in order to achieve the optimum performance.

Since the beam characteristics of a transducer are the sum of the focal characteristics of

each frequency component in the ultrasound pulse, the wider bandwidth allows more

latitude in tuning the imaging probe and controlling the frequency spectrum. By

controlling the low frequency components of the pulse, far field divergence can be

reduced, resulting in better image details in the far field.

1.6.3 Sensitivity

The sensitivity of the imaging system is a key parameter affecting the ultrasound image

quality. The sensitivity is related to the ability to detect small targets located at a specific

depth in an attenuating medium. Although it is dependent on the interaction of several

aspects of the ultrasound systems such as pulser/receiver, signal processing, display unit

21

etc., the imaging transducer plays a critical role in the system sensitivity. The transducer

variables that contribute to system sensitivity are frequency, beam geometry and energy

conversion efficiency.

The conversion efficiency of the transducer indicates how well the transducer converts

both the applied voltage into ultrasonic pressure pulse and the received echo into

electrical voltage. Sensitivity can be related to conversion efficiency. The sensitivity is

the products of transmit and receive efficiencies. The factors that influence the

sensitivity of the imaging transducer are the nature of the excitation pulse, type of

piezoelectric material, focal characteristics of the sound beam, propagating medium,

distance of the target, and the nature of the reflecting object.

The piezoelectric materials have d33 and g33 values, which are the indices of how well the

material converts the voltage signal into mechanical deformations and the mechanical

stress back into electrical voltage, respectively. Although it is advantageous for d33 and

g33 constants to be large, the commonly used materials do not necessarily have both the

highest d33 and g33 value. The other parameters to be considered are Q value, dielectric

constant, impedance and capacitance.

Although a higher degree of damping results in improved axial resolution, it lowers the

transducer sensitivity. Damping is to be selected to give a best compromise between

sensitivity and axial resolution.

Focusing is a design variable, which affects the sensitivity. Focusing regulates the width

of the sound beam. The sensitivity of the transducer changes with the variation of the

beamwidth. As beamwidth is varied by focusing, the cross sectional area over which the

ultrasound energy of the beam is distributed also varies. In the case of a non-focused

22

probe, the beam is wider and the energy content of the beam is distributed over a larger

cross sectional area. Hence the focused transducer possesses higher sensitivity. For a

given frequency and diameter, a non-focused transducer is less sensitive than a focused

one within its focal zone.

Proper selection of matching layer’s impedance and thickness will increase the efficiency

of the transducer in transferring energy during both the transmit and receive resulting in

increased sensitivity of the transducer. Poor electrical matching also results in loss in

sensitivity.

1.6.4 Backing material

When an imaging transducer is excited with an electrical voltage impulse, the

piezoelectric material oscillates at its natural resonance frequency. If there is a mismatch

in acoustic impedances of the piezoelectric material and the backing material, ringing

effect occurs during the pulse echo application, which results in lengthening of the pulse

duration causing lesser bandwidth. If absorptive backing material is used, it absorbs the

energy from the vibration of the back face of the piezoelectric material and the ringing is

suppressed. This leads to shortening the pulse duration at the expense of sensitivity,

because a large portion of the energy is absorbed by the backing material. Ideally, the

backing material should provide high attenuation and match with the acoustic impedance

of the piezoelectric material for efficient coupling. If the backing material has higher

acoustic impedance, greater sensitivity is achieved because less energy is absorbed. But

the transmitted and received pulses are longer due to increased ringing, resulting in

reduced bandwidth. The two important parameters of the backing material are the

23

acoustic impedance and the attenuation. Attenuation is the loss of acoustic energy which

is mainly due to two mechanisms namely (i) scattering loss and (ii) absorption loss.

1.6.5 Electrical matching/tuning

Usually the power delivered to the transducer by the voltage source of the system is

transmitted over a line having a characteristic impedance of 50 Ω. Electrical tuning is

often used to convert the electrical impedance of the transducer to a value closer to 50 Ω.

If the imaging transducer is not perfectly matched electrically to the pulser, then all of the

pulse energy is not delivered to the piezoelectric element. For maximum power transfer,

the input impedance of the transducer should be real and should match with the source.

A shunt inductor of the value [1/ω02C0] is used to tune out the clamped capacitance of the

transducer. But this arrangement reduces the bandwidth. Electrical matching is done by

selecting the material and tuning the probe so that there is a good electrical compatibility

between the probe and the pulsar/receiver. It involves use of single inductor, RLC

circuits and transformers. If there is no proper match between the transducer and the

pulsar, a portion of the excitation pulse is reflected back into the pulser. The ultrasound

manufacturers develop their own unique electronic circuits for electrical matching.

1.6.6 Mechanical matching (Acoustic interface/impedance characteristics)

When a sound beam travels through a material and strikes an interface or boundary, the

incident sound is partially reflected back into the first material and part will be

transmitted into the second material. The magnitude of the transmitted and reflected

signals depends mainly on the acoustic impedance of both the materials and the angle of

24

incidence of the sound beam. If the sound beam is perpendicular to the interface, the

amount of transmitted sound increases. If the acoustic impedance of the second material

is greater than the first, then more sound will be reflected at the interface. By introducing

a proper choice of material and thickness as an intervening layer, higher transmission of

sound energy from the first material through the intermediate layer and into the second

material will result as compared to the situation where there is no intermediate layer.

From the physics of wave phenomena, the optimum face thickness of the probe is one-

quarter wavelength of the frequency employed. Although three-quarter wavelength will

have the same ability to produce phase reversal and signal reinforcement phenomena,

one-quarter wavelength is normally employed to minimize the attenuation losses within

the face material. The ideal facing material impedance is the geometric mean of the

acoustic impedance of the piezoelectric material of the imaging probe and the load

material such as human tissue. ZML = MP ZZ , where ZML, ZP and ZM are the acoustic

impedances of the matching layer, piezoelectric material and the propagation medium

respectively. The quarter wavelength layer with proper acoustic impedance provides

improved transmission and reception of the ultrasound echoes.

The quarter wavelength matching layer is acting as a mechanical transformer, stepping

down the impedance change more gradually and thus reducing the degree of acoustic

impedance mismatch at the interface of the transducer and human body. At any center

frequency (fc), 100% transmission of ultrasound energy can be theoretically achieved if

the matching layer has a thickness of d = c/4fc, where c is the velocity of the sound in the

material. In practice, 100% transmission of sound is not achieved using quarter

25

wavelength matching layer, because diagnostic pulse-echo is a short, broadband pulse,

which contain a significant amount of acoustic energy, which is not at the center

frequency alone. The material characteristics of the matching layer also contribute in the

transmission efficiency. Since the single matching layer is acting as a frequency filter,

the use of matching layer reduces the bandwidth.

To alleviate the drawback of the single quarter wavelength matching layer and to achieve

improved performance, the concept of incorporating a second matching layer, which is

called multiple matching layer design is generally adopted. While there is a reduction in

bandwidth in using single quarter wavelength matching layer, one could achieve a

broader bandwidth in using two matching layer compared to the condition in using single

matching layer. The broader bandwidth leads to more energy, which results in increased

sensitivity. So the sensitivity of the transducer can be increased further by applying two

matching layers, which have acoustic impedance values between the values of

piezoelectric material and human tissue. The stepwise transition of impedances from the

transducer to the tissue allows further reduction in the impedance mismatch resulting in

decreased internal reflection and increased transmission and reception over a wide range

of frequency. This will have impact clinically in order to obtain better penetration due to

increased sensitivity, potential for using lower acoustic power output levels, and higher

frequency transducer for superior resolution.

26

1.6.7 Acoustic energy loss

There has been energy loss between the excitation pulse and the conversion of the

received echoes into electrical signal. The losses occur at many places as detailed below.

Transducer loss

The transducer cannot convert all of the applied electrical energy into acoustic waves due

to the following factors: (a) Electrical Matching: If the transducer is not electrically

matched to the pulser, then all the applied electrical energy is not delivered to the

piezoelectric material. (b) Thermal loss: A part of the pulse energy is dissipated in the

piezoelectric material in the form of heat. (c) Damping: Some of the acoustic energy is

absorbed in the backing material of the transducer. (d) Secondary modes of vibration:

Although the piezoelectric material is designed to vibrate in thickness mode, it also

oscillates in “Radial mode” which does not contribute constructively to the ultrasound

beam. (e) Mechanical matching: If the quarter wavelength matching layer is not properly

selected to act as a good mechanical transformer, the acoustic energy generated by the

piezoelectric material is not fully transmitted to the propagating medium such as human

body.

Loss in transmission medium

The transmission medium (tissue) attenuates the ultrasound at the rate of 0.5

dB/cm/MHz. The round trip attenuation loss diminishes the ultrasound energy returning

to the imaging transducer.

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Loss at the interface

When the ultrasound energy encounters an interface, the energy is not fully reflected

since part of the energy is transmitted or scattered resulting reduced energy entering the

imaging transducer.

Diffraction loss

When the ultrasound wave returns to the imaging transducer after reflecting from the

reflector, it diverges resulting less energy striking the imaging transducer.

Loss at the electrode

The existence of piezoelectrically inactive gold electrode (Zgold = 51.5 MRayls) of

thickness of about 1000 Angstrom on each side of the film for each element will

contribute some loss due to acoustic mismatch.

Conversion loss

When the echoes are received by the imaging transducer and converted into electrical

signals, similar losses as in the case of conversion of electrical voltage into ultrasound

pressure wave (due to electrical matching, mechanical matching, damping and thermal

loss) will occur.

Due to the above mentioned energy loss associated with the imaging transducer, the

sensitivity of the transducer is lowered.

1.7 Fundamentals of Array Transducer

There is a strong desire in the ultrasound community to have array transducers with broad

bandwidth and high sensitivity particularly in view of the emerging imaging modalities

such as harmonic imaging. Arrays operable at high frequency are needed for dynamic

28

focusing and to have increased frame rate for many clinical needs such as blood flow

measurements.

A typical linear array transducer consists of a long piezoelectric element strip divided

into a number of closely spaced rectangular elements. The space between the elements is

called kerf and the distance between the centers of two elements is called pitch. The size

of a pitch in a linear array element ranges from λ/2 to 3λ/2 or greater, where λ is the

wavelength in the medium of propagation [9]. Although larger pitch is not ideal, it is

possible to obtain quality images with arrays having higher pitch by using a wider

bandwidth transducer and by operating with high frame rate [10].

In a linear array system, the physical movement of the transducer is eliminated. The

deflection of the beam is done electronically. A group of elements of an array is excited

in succession so that the ultrasound beam is electronically moved across the face of the

transducer, providing an image similar to that obtained by scanning a single element

transducer manually. The microprocessor selects a “block” of elements, typically four to

eight, and pulses them. The group of selected elements simultaneously transmits the

sound waves and each wave will propagate the same distance from the transducer at a

given time. The sound waves from the elements cross each other along a line parallel to

the array resulting in formation of higher amplitude pressure waves due to constructive

interference. In the other directions, the waves are not in phase, which results in

destructive interference. Consequently, the beam is formed in the axial direction only.

The width of the elements should be larger than the wavelength of the sound in order to

29

have a highly directional beam. The resultant echoes are processed and the switching

circuit advances the “block” by one element and the process is repeated. In this manner

the entire array is sequentially activated. While the sequencing is repeated rapidly about

30 frames per second, a B scan image is obtained. Linear arrays have elements ranging

from 32 to 512 elements [9]. By electronically focusing the beam in the plane of the

scan, the position of the focus can be changed at will in the case of array transducers. By

dynamic focusing, the equipment continually adjusts the focal point to correspond with

the depth the pulse has currently reached. To focus on echoes as they return from deeper

in the body, the equipment adjusts automatically according to the distance the echo has

traveled so that on-axis echoes are always brought in phase. Because of these

advantages, array probes are commonly employed in clinical practice. The linear array

can be focused in azimuth plane only. By having a mechanical lens at the surface of the

transducer, focusing on elevation plane (perpendicular to the imaging plane) is usually

obtained, which will facilitate in determining the thickness of the slice in the imaging

plane. Due to beam broadening both at the near field and at very far field, good

resolution could not be obtained. To alleviate this problem, multidimensional arrays (2D)

are being developed [9].

1.8 Fundamentals of Ultrasonic Hydrophone Probe

Hydrophone probe is a device used to measure the amplitude of ultrasound wave

(pressure pulse) generated by the imaging transducer. This transducer converts the

mechanical energy of the pressure wave generated by the imaging transducer into an

electrical voltage signal and thus produces a voltage time waveform representative of the

30

acoustic wave pressure as a function of time. It produces an electrical voltage in response

to an applied acoustic pressure at the active element. The hydrophone probe is made

from piezoelectric polymer film called PVDF. It is the only FDA approved measurement

tool in accurately measuring the acoustic output of all the newly developed imaging

transducers before use in medical ultrasound imaging and to obtain premarket approval

[11,12]. It is the key element of the measurement system, since all the measured

parameters are directly influenced by hydrophone characteristics.

The use of an ultrasound imaging device greatly affects the patient treatment efficacy and

possible collateral tissue damage. Hence the ultrasound output from the imaging

transducer needs to be adequately characterized prior to use in clinical practice. The

hydrophone probe comprises an acoustically transparent membrane using a PVDF

polymer and having a small, central region piezoelectrically active. The hydrophone can

scan the ultrasound field to determine the acoustic field distribution, energy content,

shape and size of the beam at the focal zone. It is essential to accurately measure all the

characteristics of the acoustic waveform generated by the imaging transducer in order to

estimate the required output parameter as stipulated by US FDA and international

standard such as IEC.

The next chapter deals with background relevant to this work besides information

explaining the importance of the issue, and a review of the work done previously in this

field.

31

CHAPTER 2: BACKGROUND AND THEORY

This chapter examines the background and theory relevant to the scope of this research.

A detailed review about the existing hydrophone probe and the need for improvement are

presented. The shortcoming in the existing procedure in calibrating the hydrophone

probe is addressed along with the method proposed to resolve the issue. The existing

types of single element imaging transducers with single layer and conventional multilayer

transducers are discussed and the concept of the proposed Barker code imaging

transducer is explained. The issues involved in implementing the Barker code concept to

multielement (array) transducer are presented.

2.1 Introduction

There are two kinds of hydrophone probes in practice at present namely (i) membrane

type and (ii) needle type. The membrane type uses piezopolymer (PVDF) film, which is

stretched and glued over a supporting ring and making a small central spot

piezoelectrically active. The needle type hydrophone uses either a piezo crystal/ceramic

or piezopolymer on a tip of a needle structure (ex. hypodermic needle). Each type has

both advantages and disadvantages in application. The use of membrane type is the

acceptable device for evaluating the newly developed imaging transducers for submission

for premarket approval by FDA. However, the needle type hydrophone is acceptable for

use in the case of measuring continuous wave, since the membrane hydrophone will

introduce standing waves in such a situation. Almost all the imaging transducers use

pulsed wave.

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2.2 Ultrasonic Polymer Hydrophone Probe

Hydrophone probes are made with quartz, ceramic and polymer. Due to inadequate

bandwidth and large size, quartz and ceramic materials are not suitable for many

applications. Piezopolymer exhibits a higher maximum frequency response than the

piezoceramic.

Consequent to the discovery of strong piezoelectric properties of PVDF by Kawaii during

1969, use of PVDF in imaging transducers was reported by Ohigashi et. al and also by

Foster et. al during 1979 [13]. A new form of polymer hydrophone probe for measuring

the spatial and temporal distribution of pressure within the fields from medical ultrasound

equipment was first developed by Shotton [14] and by DeReggi and Harris [15] during

1980. The polymer needle type hydrophone probe was first developed by Lewin [16]

during 1981 and it is known worldwide as “Lewin type hydrophone”.

The hydrophone probe should be capable of capturing both the compressional and

rarefactional portion of the waveform in order to determine the peak acoustic pressure

and the integrated energy in the pulse. In order to obtain accurate information about the

temporal variation in the pulse waveform produced by the imaging transducer, the

hydrophone probe should exhibit a flat response in terms of pressure sensitivity against

the frequency over the output spectrum likely to be encountered in the frequency range

used in medical imaging. Now PVDF hydrophone probes are in widespread use

throughout the world as a reference and working device for reliable ultrasonic

measurements over a wide frequency range. Due to intrinsic broadband properties of

33

PVDF, membrane hydrophones perform linearly up to a peak pressure of 3.4 MPa [17]

and measurement of maximum pressure up to 100 MPa is possible using PVDF

hydrophone in cases like measuring the lithotriptor shock wave pulses.

The hydrophone probe operates in thickness mode with the unbacked active element,

which has acoustic impedance close to that of water. The hydrophone has a broad

frequency response due to the use of thinner films with a low Q factor. The sensitivity

varies smoothly with frequency and the hydrophones are free from reverberation effects.

Double layer hydrophone probe with thicker PVDF film has a larger reflection coefficient

than thinner hydrophones. As the film thickness decreases, the voltage sensitivity drops

because the sensitivity is proportional to the thickness. With decreasing thickness, the

thickness resonance increases in frequency but the peak sensitivity drops. The

hydrophone sensitivity depends on both the charge produced by the active element and

the overall capacitance of the device. Smaller active element will produce less charge.

Thinner polymer film would have lesser pressure sensitivity since the active volume of

piezoelectric material would be smaller.

In the field of ultrasound, both the hydrophone probe and imaging transducer always go

hand in hand. One cannot be of any use without the other. Both are part and parcel of

acoustic measurement system. Hydrophone probe is used in the measurement of acoustic

fields produced by the imaging transducers.

34

2.2.1 Need for wider bandwidth hydrophone probe

It is known that the polymer hydrophone probe is the only FDA approved, standard

measurement tool in accurately measuring the acoustic output of all the newly developed

imaging transducers before use in medical ultrasound imaging [12]. There is an

increasing demand for precise measurements to be performed on the acoustic fields

emitted by medical imaging transducers. The finite amplitude distortion is a common

characteristic of large amplitude pressure waves, and it is especially prevalent in focused

ultrasonic fields. Because of finite amplitude effects in water or tissue, diagnostic

ultrasound pressure pulses can have significant high and low frequency spectral content.

To record these pressure pulses accurately, the hydrophone probe must have sufficient

broadband response [18]. The two types of double layer hydrophone probes commonly

used have frequency response up to 20 and 50 MHz only. Harris [18] has pointed out the

need for hydrophones with much broader bandwidth in order to measure the center

frequency of the imaging transducer and higher harmonics generated through nonlinear

propagation effects in water. The pulse parameters determined from the data obtained

from the hydrophones with inadequate bandwidth showed large errors. The pressure

pulse and the hydrophone are characterized by the pulse center frequency fc and the

hydrophone thickness resonance frequency fh respectively. Harris has also pointed out

that the ratio of fh/fc should be greater than about 10 on order to keep the error below

±5% in calculating the pulse average intensity [18]. IEC recommends that the bandwidth

of the hydrophone/amplifier combination needs to be greater than 8 times that of the

center frequency of the pulse [IEC 87(Co) 6, 1988]. As per FDA 510(k) [11] and AIUM

[12] guidelines, it is desirable to have the ±3 dB bandwidth of the hydrophone probe wide

35

enough to make the acoustic measurements at the upper frequency limit of 8fc in order to

have accurate measurements of high frequency/wide bandwidth transducers exhibiting

nonlinear distortion.

Typical double layer hydrophone probes are customarily made with either 25+25 micron

thick or 9+9 micron thick PVDF film, which have the measured thickness resonance of

about 20 and 50 MHz respectively. Hence an improved design of hydrophone probe with

enhanced bandwidth capable of characterizing the imaging transducers emitting high

frequency components was considered necessary. Therefore, besides fabricating and

evaluating the conventional double layer polymer hydrophone probes with similar

thickness of PVDF film (25+25 µm & 9+9 µm), polymer hydrophones using dissimilar

thickness (25+9 µm) were developed to maximize the bandwidth and sensitivity. The

hydrophones were characterized for acoustic sensitivity, frequency response, bandwidth,

angular response, effective aperture size and orientation effects. The literature search

revealed that the development of improved design of double layer hydrophone probe to

enhance the bandwidth and sensitivity using dissimilar thickness of (25+9 µm) PVDF

film was not published by any one previously.

2.3 Calibration Techniques

There are mainly two types of calibration techniques namely (i) primary calibration

technique and (b) secondary calibration technique. In the primary calibration, there is a

distinction whether the calibration technique used is a primary technique or the technique

uses a primary standard device. The primary technique uses independent measurement of

36

certain parameters, which are directly traceable to primary standards. The primary

calibration techniques are (i) optical interferometry, which is traceable to the

measurement of length [19], (ii) radiation force balance technique, which is traceable to

the measurement of force [20], (iii) calorimetry, which is traceable to the measurement of

temperature [21], (iv) planar scanning technique, which is traceable to radiation force [1],

and (v) reciprocity [22]. The secondary calibration technique uses the substitution

method, which compares the unknown value of the hydrophone to be calibrated with a

known value of a reference hydrophone that has been calibrated previously. So, the

substitution technique requires the use of a reference device. The calibration technique

used in this work namely “planar scanning technique” requires a primary standard

acoustic source traceable to the radiation force approach. Hence the calibration

performed by using the planar scanning technique is an absolute calibration.

The output of diagnostic ultrasound equipment is regulated by the Food and Drug

Administration (FDA) in United States [11] and worldwide, by IEC standards [23].

Manufacturers of ultrasound equipment, when submitting application for 510(k) process,

need to provide all relevant acoustic output data including the maximum acoustic

pressure amplitude and pressure distribution produced by the ultrasonic device to be

marketed. However, the absolute calibration data for such hydrophones are available

usually at frequencies above 1 MHz. The existing IEC standard for primary calibration

of medical ultrasonic hydrophones covers the frequency range from 0.5 MHz to 15 MHz

[23] and a standard concerned with frequency range from 15-40 MHz is under

development. Harris has pointed out that the knowledge of the frequency response of the

37

hydrophone probes below 1 MHz is of importance and has provided evidence that

inadequate low frequency response of the hydrophone/preamplifier assembly introduces

distortions in the measured pressure-time waveforms generated by ultrasound diagnostic

devices [24]. He also pointed out that, consequently, a significant measurement error

(exceeding 30%) could be introduced in determining values of mechanical (MI) [24]. In

the following publication Harris showed that the hydrophone bandwidth must extend to at

least 10 times below the diagnostic pulse center frequency to minimize the error in the

measurement of peak rarefactional pressure amplitude and thus Mechanical Index (see

Eq.2.1, below) to approximately 5% [25]. It might be appropriate to briefly note that

Mechanical Index (MI) is defined based on a theoretical assessment of the likelihood for

cavitation to occur in a medium containing cavitation nuclei with a wide size distribution.

Its value can be determined as [26]:

MI = 10 f015.0f

cr

c

zp − (2.1)

where pr is the peak rarefactional pressure amplitude in MPa, fc is the center frequency in

MHz of the acoustic source measured in water at the axial distance z in cm where the

derated pulse intensity integral (i.e. the integral of the pressure squared) exhibits a

maximum [26]. The derating factor 10 f015.0 cz− is intended to account for attenuation in

tissue by using an estimated average attenuation of 0.3 dB/cm.MHz [12]. As the portion

of the pressure waveform where pr occurs is dominated by the low frequency

components, it is essential that the frequency response of a hydrophone probe below 1

MHz be known. More specifically, due to the nonlinear propagation effects in water and

with diagnostic imaging transducers operating frequently at center frequencies of 2 MHz

and below for harmonic imaging, the spectrum of pr may contain frequencies down to

38

100 kHz [24]. Also, a number of ultrasound devices operate below 1 MHz and their

acoustic output needs to be measured or monitored. These devices include lithotripters

operating at center frequencies between 200-1000 kHz, nebulizers, ultrasound-healing

devices which utilize frequencies ranging from 30-600 kHz, and High Intensity Focused

Ultrasound (HIFU) systems which are designed with center frequencies between 400-750

kHz. Also, the commercially available ultrasonometer (Lunar) used for evaluating the

risk of osteoporotic fracture in postmenopausal women operates at 500 kHz. In addition,

many of the therapeutic ultrasound units operate at a center frequency of 1 MHz and

below. Hence, there is a well-defined need to have a dependable method for absolute

calibration of hydrophones in the frequency range 0.1–1 MHz.

As already noted, at present, no commercially available hydrophone probes are available

with sensitivity information below 1 MHz. Harris has examined the behavior of

hydrophones from 0.2 MHz – 2 MHz using broadband pulse technique [27]. He

observed that the frequency response of the PVDF membrane hydrophones was

essentially flat below 1 MHz and his results were corroborated using the swept frequency

technique combined with reciprocity technique [28]. In this work the applicability of

planar scanning technique was examined as an alternative method for low frequency

calibration of the probes. While measurements using planar scanning technique have

been reported in the megahertz range of frequencies [1, 29, 30], it appears that this

technique was not explored at frequencies below 1 MHz.

A brief description of the two methods used in this work is given below:

39

2.3.1 Planar scanning technique

In the planar scanning technique, the total acoustic power produced by the source was

determined by using the hydrophone to be calibrated by performing raster scan in the far

field of the acoustic source and then carrying out the spatial integration of the square of

the voltage measured at the hydrophone’s terminal. By using the Radiation Force

Balance (RFB), the acoustic output of the source was measured. By comparing the

power obtained from the raster scan measurements with the power measured with the

radiation force balance (RFB), the hydrophone’s sensitivity was determined [31]. This

method provides hydrophone’s sensitivity at discrete frequencies.

2.3.2 Time Delay Spectrometry technique combined with substitution method

To mention briefly, the Time Delay Spectrometry (TDS) method [2, 3, 4], which is swept

frequency (rather than a single frequency) technique works in the following way: The

wideband source transducer was driven by the swept sine wave signal from the tracking

generator of the HP3585A spectrum analyzer with a built-in frequency offset unit and

amplified by a linear broadband power amplifier. The received signal from the

hydrophone was fed into the band pass filter of the spectrum analyzer, which is swept

with appropriate time delay in relation to the transmitter-driving signal. The signal

corresponding to the direct propagation delay between the transmitter and the receiver

alone is picked up after filtering out the other spurious signals. Briefly mentioning the

substitution calibration method, it is performed in the following way: The frequency

response of the given hydrophone/source combination was compared to that obtained

from the reference hydrophone used with the same source transducer. First the reference

40

hydrophone is placed in the acoustic field at the far field of the source transducer and the

acoustic signal is maximized and captured into the spectrum analyzer. The reference

hydrophone is removed and the hydrophone to be calibrated is replaced at the same

location; the signal is again maximized and captured into the spectrum analyzer. The

difference between the two spectra is added to the known frequency response of the

reference hydrophone to obtain the frequency response of the hydrophone to be

calibrated. The advantage of TDS technique is that it allows the end-of-cable receiving

voltage sensitivity to be determined virtually as a continuous function of frequency. This

feature effectively eliminates the possibility of overlooking rapid sensitivity variation of

hydrophones.

2.3.3 Need for developing a new measurement procedure

The acoustic output parameters of all the newly developed ultrasonic imaging transducers

are to be measured before use in clinical practice. The Mechanical Index (MI), a

predictor of potential bioeffects in ultrasound imaging is displayed on the ultrasound

equipment, which gives an indication of the potential for mechanical damage to exposed

tissues. AIUM standard stipulates that a hydrophone with a lower frequency limit of

fc/20 will provide more accurate measurement of the peak rarefactional pressure, and thus

enable to determine the Mechanical Index more accurately [12, 24]. Inadequate low

frequency response of the hydrophone introduces error in the measurement of peak

rarefactional pressure amplitude, particularly below 1 MHz [32]. One of the problems

associated with assuring an adequate low frequency response for hydrophones is the lack

of available calibration techniques below 1 MHz. For example, for adequate

41

characterization of acoustic field produced by an imaging transducer operating at the

center frequency of 10 MHz, the sensitivity of the hydrophone at frequencies in the

vicinity of 500 kHz should have been known. Because the portion of the pressure

waveform where pr occurs is dominated by the low frequency components, it is essential

that the frequency response of all newly developed hydrophone probes below 1 MHz be

known.

The absolute calibration data for such hydrophones are available usually at frequencies

above 1 MHz. No national or international standard exists describing the procedure for

absolute calibration of the ultrasonic hydrophone probes at frequencies below 500 kHz.

Hence there is a well-defined need to have a dependable method for absolute calibration

of hydrophone probes in the low frequency range (below 1 MHz). Hence, in order to fill

the gap in the knowledge base, the applicability of planar scanning technique was

examined in this work as an alternative method for low frequency calibration of

hydrophone probes. Although measurements using the planar scanning technique have

been reported in the megahertz range of frequency, this technique has not been explored

at frequencies below 1 MHz

2.4 Single Element Imaging Transducer

2.4.1 Transducer requirements

Imaging transducer requirements mainly include high sensitivity and exceptionally broad

bandwidth to cover the frequency range of interest with a single transducer.

42

2.4.2 Conventional piezoceramic imaging transducers

The conventional imaging transducer used in diagnostic systems is limited due to the

inherently strong resonance of the piezoelectric ceramic material. To optimize the image

resolution and penetration depth, multiple transducers operating at different resonance

frequencies are being used at present to cover the frequency range of interest during

diagnosis due to the limited bandwidth of the PZT (Lead Zirconate Titanate) ceramic.

The PZT transducers consist of thin disks of ceramics, operated at resonance and the

bandwidth is determined by the thickness resonant frequency (λ/4 or λ/2) of the ceramic.

Because of the brittleness of the PZT material, thin disks are difficult to fabricate.

2.4.3 Conventional polymer imaging transducers

A brief outline of conventional polymer transducer using single layer and multiple layers

is given below in order to have an understanding before presenting the Barker code

imaging transducer. The use of PVDF in medical imaging was first reported by Ohigashi

and Foster during 1979 [32].

2.4.3.1 Single element, single layer transducer

The single element transducer needs to be moved physically over the area to be examined

and a two-dimensional image (B Scan) is formed by combining the lines of information

generated. The beam is focused using a fixed lens, which is simple and reliable but has

the major drawback that the focal point cannot be adjusted. The limitation of the single

layer imaging transducer is that when operated with short impulse, only small acoustic

energy can be generated, which limits penetration depth.

43

2.4.3.2 Single element, multilayer polymer transducer

The weaknesses of PVDF having lower electromechanical coupling factor (kt = 0.20), and

dielectric constant (εS = 6) compared to PZT can be minimized by the use of multiple

layers [33]. The multilayer structure is an alternative technique to deliver more acoustic

output power from a voltage source drive. In a conventional multilayer transducer, N

identical PVDF films of thickness l are stacked, while the films of alternating layers have

reversed polarity. All layers have the same lateral dimensions, which are large compared

to the thickness of the layer. The layers are mechanically connected face to face resulting

electrically in parallel and mechanically in series. The transducer is resonant at the

desired operating frequency where the total stack thickness is λ/4 or λ/2 wavelength.

Since the layers are acoustically in series, the resonance frequency is inversely

proportional to the total thickness of the stack:

ω0 α Nl1 (2.2)

where l is the thickness of the layer and N is the number of layers. So by increasing the

thickness of the stack, the resonance frequency is downshifted.

In the transmit mode, multilayer structure gives increased acoustic output for a given

source voltage. Since each layer is sandwiched between ground electrode and a signal

electrode, the electric field direction is constant with the poling direction when the

voltage is applied to the multilayer structure. Therefore, the layers expand or compress

uniformly. Since all layers move together, the output pressure is increased proportionally

with the number of layers. The electrical power input is

44

Pin α elZ

V 2

, Zel α N1 (2.3)

The advantage of N layers is to decrease the electrical impedance by a factor of N. For a

particular stack thickness Nl, and source voltage input V, the total voltage induced across

the stack is the result of effective voltage of NV. The acoustic output power is

proportional to the square of the pressure amplitude, which leads to:

Pout α N2 (2.4)

2.4.4 Barker code polymer imaging transducer

The inherent drawbacks in the conventional multilayer transducer can be overcome by

the proposed design, which utilizes the concept of Barker code [34]. Prior to the

presentation of the principle of operation of Barker code multilayer polymer transducer, a

brief description of Barker code concept is outlined below.

2.4.4.1 Barker code concept

Barker codes are binary codes named after R.H. Barker, who originally developed the

codes to synchronize signals in digital communication systems. Barker code is used for

the internal arrangement of polarization pattern of piezoelectric layers, in which pulse

compression is effected [35]. This property was used to implement the pulse-echo Barker

code transducer. When a Barker code transducer is excited, the transmitted acoustic

wave is the superposition of the acoustic wave generated by each piezoelectric layer in

the assembly. While in receive, the pressure wave echoing from the reflection and

entering into the transducer is the Barker coded pulse sequence transmitted by the same

transducer. The signal compression is brought about by receiving these pulse trains by

45

the same transducer. With this method, the sensitivity of PVDF transducer is raised by

pulse compression.

2.4.4.2 Principle of operation

The schematic representation of the Barker code arrangement is shown in Figure 2. In

this design, identical elements of thickness l, electroded on both sides are stacked in such

a way that the polarization pattern is chosen according to Barker code, and are operated

in their thickness mode. [35, 36]. The layers are glued on to the plane surface of an

acoustically matched and absorbing material to make an ultrasonic multilayer transducer

having total thickness of Nl, whose frequency response essentially corresponds to that of

a transducer made of a single layer of thickness l [36]. The transmit/receive switches are

provided in such a way that the layers are connected electrically in parallel in the

AcousticBacking Re

flect

or

Piez

oelec

tric

Laye

r

Oscilloscope

p(x)

x0 3l

Medium of propagation+ - + -- +

V(t)

t0

3

-1 -1

3l/c 6l/c

Figure 2: Schematic representation of the Barker code arrangement

46

transmitting mode and electrically in series in the receive mode. When such a transducer

is excited, each individual film generates a pressure pulse where the polarities

(compression or tension) correspond to the direction of the polarization and to the sign of

the applied voltage. Hence a train of pressure pulses with Barker coded structure is

produced in the medium of propagation [36]. At a particular time, each of the individual

pressure pulses sent out from each layer will reflect back and will reach its corresponding

layer. When this occurs, the voltage generated in each layer will exhibit the same

polarity, thus the output voltage produced at the terminal of the transducer will be N

times larger than the voltage generated in a single layer. During the rest of the time, the

voltage generated in each layer exhibits different polarity and some of the voltages will

be cancelled out [37]. Hence the stress wave echoing from the reflector and entering into

the transducer is the Barker coded pulse sequence transmitted by the same transducer.

The resulting signal compression gives an electrical output signal consisting of a

dominant sharp peak and some ripples due to the general properties of the Barker code

[35]. The maximum receive sensitivity gain (the peak output voltage) compared to a

single layer transducer is a factor of N, which corresponds to the Barker code length of N.

The other feature of the Barker code is that the peak side lobes are less than or equal to a

factor of one in magnitude. The ratio of the amplitude of the main peak and the side

lobes is N:1 [36]. The side lobes, which contain the minimum energy, are uniformly

distributed. Moreover, the higher receive voltage minimizes the effect of preamplifier

noise and the diagnostician can select a lower receive gain setting in the system, which

results in increased signal-to-noise ratio (SNR).

47

In contrast to a conventional piezoceramic [38] and composite material transducer in

which the bandwidth is determined by the thickness resonance frequency of the active

element and operates at resonance, the piezopolymer transducer is designed to operate at

frequencies below the resonance [39]. In the very near field, at which the distance is very

small compared to the radius of the multilayer transducer, the received signal should be

proportional to the autocorrelation function of the Barker coded pulses. As distance

increases, the time-derivative of the autocorrelation function would be received [35]. In

the Barker code multilayer approach, bandwidth is determined by the thickness of the

single PVDF film and the sensitivity is determined by their overall thickness. Hence, the

pulse-echo sensitivity is increased compared to single layer structure while retaining

inherent wideband characteristic of PVDF film.

The Barker code concept was originally introduced to the single element transducers by

Sung in 1984 [35] and Platte in 1987 [36]. Zhang [39] had extended their work by

carrying out the modeling and conducting performance evaluation of the Barker code

transducers in transmit-receive mode. Zhang’s thesis work covered on the experimental

data using two Barker code transducers, one as a transmitter and another as a receiver. In

contrast, this research work focused on implementing the Barker code concept in “pulse-

echo” mode, both in the case of single element and multielement (array) polymer

transducer, which is the practical application in ultrasound imaging. The development of

Barker coded polymer array probe in this research work for medical ultrasound

application was the first one of this kind and no one has developed this type so far.

48

2.5 Multielement (array) Imaging Transducer

The lateral resolution at any plane is limited by the depth of focus. The depth of focus is

defined as that region where the geometric shadow of the rays is equal to the diffraction

limited size and is given by [40]:

df

fd λ=

∆ i.e. ∆ = 2

2

dfλ ∴ 2

222d

fλ=∆ (2.5)

where ∆ is the distance on either side of focus and thus 2∆ is the total depth of focus.

The width of the main central lobe is determined by the width (X) of the array. If the

width of the element (W) is an appreciable percentage of the combined size of width of

the element and the kerf (S), then the side lobes are automatically attenuated

significantly. By using large number of elements, where X/S is a relatively large number,

the side lobes can be reduced but at the expense of increased complexity [40]. The array

can be focused in the near field to obtain the diffraction limited resolution by using a lens

in the X, Y plane and the focal length in such case is given by [40]:

f = ( )21 nnR−

(2.6)

where, R is the radius of curvature of the lens surface and (n1-n2) is the difference in the

refractive index between the lens material and its surround. The lens is required because

electronic focusing focuses the beam in the direction perpendicular to the elements only.

The resolution of an array is usually defined as equal to λ fn, where fn is the focal number

defined as Z/D, while D is the total length of the array (aperture) and Z is the focal

distance.

49

2.5.1 Side lobes and grating lobes

The side lobes are present next to the main lobe of the response. If the diameter of a

single element transducer is large, a directional beam will be formed due to

predominance of the constructive interference effect. There are weak areas of

constructive interference in some other directions at certain angle to the main beam

resulting in low intensity subsidiary beams, which are called side lobes. In the case of

array transducers, they are called grating lobes, which appear due to the discreetness of

the array. The side lobes and grating lobes introduce uncertainty in the origin of the

echoes. The ultrasound image is based on the display of echoes from the reflecting

structures due to the main beam only, which should be narrow both in transmission and

reception. The side lobes which occur at certain angle due to constructive interference in

view of the regular spacing of the element is called grating lobes, which is given by [9]:

= −

gn

gλφ 1sin (2.7)

where n is an integer number 1,2,3… , g is the pitch and gφ is the angle at which the

grating lobe occurs. For the grating lobes to occur at angles greater than 900, the pitch of

the array (g) should be smaller than λ/2. The amplitude of the grating lobe is determined

by the width of the element. If the width is small, the amplitude of the grating lobe will

be larger compared to the main lobe. The angle of grating lobe and the main beam is

given by:

sinϕ = sinθ + pλ (2.8)

where ϕ is the angle of grating lobe, θ is the angle of main lobe, λ is the wavelength and

p is the pitch. The effect of side lobes will be the distortion of the image. The side lobes

50

are located approximately at a distance 1.5λZ/D from the center of the response, where Z

is the focal length and D is the aperture length. The grating lobes are spaced in the x

direction by a distance of λZ/d, where d is the center-to-center spacing of the elements.

2.5.2 Cross-talk effects

The cross-talk effect is an important aspect in evaluating the performance of the array

transducer, because it affects the ring down through delayed signals from the adjacent

elements and angular dispersion by increasing the effective element size. In a linear

array transducer, the voltage generated by the receiving element should be related to the

incident pressure wave on the receiving elements only or to the excitation voltage applied

to the corresponding elements only. But in practice, if a single or a group of elements are

fired during transmit mode, not only the specified elements generate an acoustic signal

but also it causes the adjacent elements to generate the acoustic signal. In the same way,

the echoed pressure pulses from the reflecting target, not only causes generation of

electrical voltage from the respective transmit elements, but also some additional signal

from the adjacent elements. This phenomenon of generation of additional pressure

wave/voltage unconnected with the excitation events of the elements of an array is called

cross talk effect and the mechanism of the cross talk is usually classified mainly as

electrical and acoustical. The cross talk effect will arise due to interfering waves

propagating along the inter element. The cross talk leads to loss in sensitivity due to

large parasitic capacitance within the array structure. If the cross talk effect is

sufficiently strong, then the directivity pattern of the individual element will differ from

the ideal one.

51

2.5.2.1 Electrical cross talk

The electrical cross talk is related to the energy coupling between adjacent array

elements, which may be due to the generation of fringe electric fields along the edge of

the electrodes and influencing the neighboring elements. The cross talk may also be

caused by parasitic capacitance between the adjacent elements.

2.5.2.2 Acoustical cross talk

Acoustical cross talk arises due to lateral acoustical coupling among the adjacent

elements by the surface wave and Lamb waves. If a longitudinal wave reaches the

surface of a medium with an angle of incident other than normal, then a surface acoustic

wave is generated and it will propagate along the surface of the medium. The resonance

frequency is determined by the periodic loading of the surface by the array elements. So,

when the reflected pulse wave is incident on the surface of an element in an array, the

surface wave will be generated and travels to the adjacent elements. A Lamb wave will

be generated if the incident wave hits a thin layer between two other layers, and it will

propagate along the thin layer. The Lamb wave causes cross coupling between adjacent

elements similar to surface wave. The surface wave and Lamb wave emerge due to

parasitic modes of vibration and cause additional energy dissipation in an array.

52

2.5.2.3 Electromechanical coupling

Since the array elements are in the same sheet of PVDF, the mechanical vibration caused

by the incident pressure wave on a particular element may be coupled to the neighboring

elements, which will cause output voltage generated across these elements. Similarly, in

transmit mode, the applied voltage on a particular element not only generates longitudinal

strain in the thickness direction but also in the transverse direction due to transverse

piezoelectric coupling. This vibration may induce voltages in the neighboring elements.

Cross talk effects in imaging transducers are not desirable. They introduce a phase shift

to the array element causing loss of phase coherence, resulting in increased insertion loss

and loss in resolution affecting the image quality.

The next chapter deals with the critical analysis of material properties and suitability of

piezoelectric materials for the use in imaging transducers.

53

CHAPTER 3: PIEZOELECTRIC MATERIALS

A brief description of different properties of various piezoelectric materials used in

imaging transducers are presented here in order to arrive at a decision about the

suitability of using a particular material for the imaging transducer.

3.1 Introduction

The phenomenon that the application of voltage on a piezoelectric material makes

changes in its physical dimension and the application of pressure on a piezoelectric

material produces electric field is called pressure-electric effect or piezoelectric effect.

The piezoelectric effect was first discovered by Pierre and Jacques, French Physicists

during 1880 [9].

The piezoelectric material consists of randomly placed numerous electric dipoles. By

polarizing the piezoelectric materials, the dipoles are aligned uniformly in the direction of

polarization resulting in change in the thickness of the material. One method of

polarization is carried out by applying a strong electric field along the direction in which

the piezoelectric effect is preferred and simultaneously heating it just above its Curie

temperature and then cooling it slowly to the room temperature by still keeping the

applied electric field.

3.2 Material Properties

The piezoelectric material has several important properties, which will have impact on

the performance of the imaging transducer. A brief description of some of the relevant

54

properties is given below in order to have an understanding as to how each parameter

affects the performance of the imaging transducer.

3.2.1 Piezoelectric stress constant (e33)

The stress produced in a piezoelectric material upon the application of a unit electric

field, while being clamped, is called the piezoelectric stress constant e. The unit is

Newton/V-m or Coulombs/m2. The larger the magnitude of the stress constant, the

greater the coupling between elastic and electrical effects.

3.2.2 Piezoelectric strain constant (d33)

The piezoelectric crystal is defined in x, y, z axis or direction and indicated by 1, 2, 3.

The polarization direction is the one in which a stress produces an electric field. The

piezoelectric crystal when it is cut with its surface perpendicular to x-axis, is called x cut

and so on. The z direction is normally used to indicate the polarization direction of a

crystal. The strain produced in the z direction while applying the electric field with no

external stress is called piezoelectric strain constant, d33.

3.2.3 Transmitting constant

The strain produced per unit of applied electric field with no external stress is known as

the transmitting constant, d. The unit is Coulombs/newtons or meters/volt. The

relationship between stress constant and transmitting constant is e = cE.d, where cE is the

elastic constant of the material.

55

3.2.4 Receiving constant

For a unit applied stress, the open circuit electric field produced is called the receiving

constant, g. The unit is V-m/Newton. This relates the electric field to the stress as gh =

dh/εT, where dh is the hydrostatic piezoelectric coefficient and εT is the free dielectric

constant.

3.2.5 Dielectric permittivity (constant)

The dielectric constant determines the electrical impedance of the piezoelectric material.

A large dielectric constant is important for enabling a good electrical impedance match to

the system electronics. While the piezoelectric material is clamped in such a way that it

cannot move while applying a voltage or if there is no strain, the measured dielectric

constant is denoted as “clamped dielectric constant”, εS. While the piezoelectric material

is kept in such a way as to move freely without any restriction, the measured dielectric

constant is denoted as “free dielectric constant”, εT. The relationship between the

transmitting and receiving constant is defined as εT = d/g. If the dielectric constant of a

piezoelectric material is low, then the input electrical impedance will be high,

necessitating higher drive voltage to generate the required output acoustic power.

3.2.6 Electromechanical coupling coefficient

The electromechanical coupling factor determines the efficiency of the piezoelectric

material in converting the mechanical energy into electrical and vice versa in a single

thermodynamic cycle. A failure to convert a large fraction of energy in a single cycle

will result in loss of sensitivity and loss of bandwidth if the energy is converted in a later

56

cycle resulting in spreading the response time. High coupling factor leads to improved

axial resolution, broader bandwidth and higher sensitivity. The electromechanical

coupling coefficient (ECC), which is a measure of the performance of a material as a

transducer is expressed as [9].

ECC = 1−S

T

εε (3.1)

The piezoelectric material can have resonance in both lateral and thickness modes

depending upon its dimension. The “thickness mode ECC” is designated as kt, which is

the measure of the ratio of the transferred electromechanical energy in the thickness

mode to the total input energy. This electromechanical coupling factor (kt), which is a

dimensionless quantity, is one of the important factors in obtaining higher acoustic output

power, which gives the better measure of the acoustic radiating power of the transducer.

The pulse echo response of an imaging transducer depends on kt, while larger kt

contributes to the increase in response amplitude. The larger the value of kt, the higher

the piezoelectric coupling between the acoustic and electrical properties of the material.

3.2.7 Dissipation factors

3.2.7.1 Mechanical loss tangent

The larger mechanical loss decreases the response time resulting in improvement in the

sharpness of the response, although it reduces the response amplitude considerably. The

mechanical quality factor Qm is the reciprocal of the mechanical loss tangent; i.e. Qm =

1/tanδm.

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3.2.7.2 Electrical loss tangent

With increase in electrical loss tangent, conversion loss will increase due to dissipation of

energy within the transducer. The dielectric loss will not affect the response time but the

amplitude will be decreased slightly. The electrical quality factor is the reciprocal of the

electrical loss tangent i.e. Qe = 1/tanδe. A minimal electrical loss tangent reduces the

amount of heat generated by the material on transmit and improves the signal to noise

ratio during receive mode.

3.2.8 Acoustic impedance

The acoustic impedance matching of the piezoelectric material of the imaging transducer

matching with that of the propagation medium such as water or tissue is a very important

and sensitive factor on the response time of the waveform. The response time is strongly

dependent on Z. When operating into water or tissue, low impedance is a distinct

advantage as it results in relatively low reflections at the interface of the transducer-

water/tissue. This will lead to broad bandwidth without requiring matching layers.

3.3 Piezoelectric Materials for Imaging Transducers

3.3.1 Piezoceramic

Although piezoelectricity was discovered in 1880, the fact that piezoelectricity could be

observed in ceramics was not published until 1946. The piezoelectric properties were

discovered in barium titanate in 1952. A family of ceramics made from lead zirconate

and lead titanate are collectively known as PZT.

58

Lead Zirconate Titanate (PZT) has been used in medical ultrasonic transducer design

since the 1970s. The advantage of PZT is the high electromechanical coupling

coefficient in thickness mode (kt = 0.50), high relative dielectric constant (600) and low

mechanical loss tangent (tanδm = 0.004), and dielectric loss tangent (tanδe = 0.002). For

medical ultrasound transducers, the electromechanical coupling factor kt is one of the

most important factors, which characterizes the energy conversion in the thickness mode

vibration, resulting in large acoustic power. Since the pulse echo response amplitude

strongly depends on kt, the larger kt contributes to the increase in the response amplitude

to a great extent. The major disadvantage of PZT is its brittleness. Ceramic transducers

have struggled with reproducibility, and fabrication difficulties. Because of the high

acoustic impedance of piezo-ceramic (34 MRayls) compared to human tissue (1.7

MRayls), there is an acoustic mismatch. Because of this mismatch, acoustic waves

leaving the ceramic and entering the water/tissue are highly reflected causing

reverberations in the ceramic leading to very sharp resonant peak, resulting in narrow

bandwidth. It also leads to an impulse response ringing for several cycles, and inefficient

power transfer. To improve these characteristics and increase the efficiency of the

transducer, matching layers having intermediate impedances need to be placed between

the ceramic and water/tissue, which further limits the bandwidth. The typical bandwidth

attainable with PZT is about 60%. PZT could not be used in imaging transducer above

15 MHz due to difficulties in making and working with the very thin plate, which is very

brittle.

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3.3.2 Piezocomposite

By combining the piezoelectric ceramic with a piezoelectrically inert polymer, many

composite configurations are made, which provide new properties that could not be

realized with either one alone. There are three types of piezocomposite that are

commonly used in imaging transducers. They are (a) 1-3 piezocomposite (b) 2-2

piezocomposite (c) 0-3 piezocomposite.

In a piezocomposite, one component is piezoelectrically active material and the other is

an inert (insulator) material. The term connectivity means the number of directions

through which the material is continuous. The X, Y, Z axes are also called 1, 2, 3

directions. If the ceramic connectivity is 1, and the polymer connectivity is 3, then the

resulting material is called 1-3 composite

The 1-3 piezocomposite consists of arrays of piezoelectric ceramic rods arranged in a

polymer matrix. The 2-2 composite have arrays of piezoelectric ceramic strips separated

by inert polymer strips or air. The 0-3 piezocomposites have piezoelectric ceramic

particulates dispersed in a polymer matrix.

The low-density polymer assures that the piezoceramic has a high electromechanical

coupling coefficient. Since the density and velocity of sound of the composite is reduced,

the acoustic impedance of the piezocomposite is lower compared to ceramic. In a

piezocomposite, only the thickness mode is excited in order to obtain the compact

60

impulse response. The composite material can be made flexible, having

electromechanical coupling coefficient higher than PZT.

A mixture of polymer material and PZT (1-3 composite) is used to modify specific

feature of the ceramic for applications in the 1 to 7.5 MHz range [9]. These 1-3

composites, consisting of small PZT rods embedded in a low density polymer having a

lower acoustic impedance (4 to 25 MRayls) than conventional PZT, better match the

acoustic impedance of human skin. Besides low acoustic impedance, the composite

material has the advantage of mechanical flexibility, improved electro-acoustic efficiency

with an electromechanical coupling coefficient higher than that of PZT. The piezo-

composite makes a short pulse with increase in amplitude compared to PZT resulting in

increased bandwidth. The typical bandwidth attainable with piezo composite is about

80% [33].

3.3.3 Single crystal

Ever since the discovery of the piezoelectric effect in barium titanate in 1952, only the

ceramics are used for the piezoelectric activity, primarily due to high coupling factor of

PZT ceramic (> 0.75) over any known crystal. Besides, the single crystal of usable size

could not be grown until 1982, when Kuwata et. al discovered high piezoelectric

coupling (> 0.9) in a solid solution of lead zirconate niobate (PZN) and lead titanate (PT)

known as PZN/PT [41].

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Lithium Niobate (LiNbO3)

A typical piezocrystal is Lithium Niobate (LiNbO3), which is one type of single crystal,

which has a high thickness mode coupling coefficient (kt = 0.49) and low relative

clamped dielectric permittivity (εS/ε0 = 28) [42]. The low relative permittivity leads to

increase in electrical impedance. The high acoustic impedance of LiNbO3 (34 MRayls)

poses problems as PZT in acoustic impedance matching. The measured –6dB bandwidth

varied from 50% to 70% [42].

Potassium Niobate (KNbO3)

Potassium Niobate (KNbO3) possesses a low dielectric constant, and high coupling

coefficient kt = 0.68). The high coupling factor leads to improved axial resolution and

high sensitivity. A bandwidth of 64% was achieved [43].

PZN/PT and PMN/PT Single Crystal

Single crystals are made from the solid solution of Pb(Zn1/3 Nb2/3)O3 abbreviated as PZN

and PbTiO3 abbreviated as PT. The PbTiO3 is a typical piezocrystal. Kuwata et.al has

reported a very high electromechanical coupling coefficient (k33 = 0.92) in PZT/PT single

crystal and the imaging probe fabricated using PZN/PT 1-3 composite revealed a

bandwidth ranging from 74 to 141% [7]. Since the sound velocity of PZN/PT is less than

70% of PZT, PZN/PT offers lower acoustic impedance.

Researchers have also developed a single crystal with 0.91Pb(Zn1/3 Nb2/3)O3 –

0.09PbTiO3 [41].

62

There are several issues in using single crystals for imaging transducers. There is a

practical difficulty in growing the crystal without defects and consistent properties within

a piece, within a batch and from batch to batch. The other issues are maintaining proper

crystal orientation, avoiding chipping and cracking, achieving good electrode and

adhesion, problems of depoling and repoling etc. Although these materials can transmit a

large amount of energy, they cannot transmit high bipolar voltage or voltages with

incorrect polarity. Once the above practical problems are resolved, single crystals may

dominate the ultrasound industry due to their capability of giving simultaneous broad

bandwidth and high sensitivity, which is the prime demand of the ultrasound world.

3.3.4 Piezopolymer

PVDF is a semicrystalline polymer. It was patented by Ford and Hanford in 1948 [44].

In 1969, Kawai reported the discovery of a large remnant polarization in oriented films of

PVDF [44]. PVDF exhibits non-polarizable form after cooling from the melt. In order to

make it piezoelectric, it is to be converted into polarizable form by stretching the film to

approximately 4 to 5 times its lateral dimensions. The lateral stretching induces a high

degree of molecular orientation in the direction of stretching. The films can be stretched

in both lateral direction called biaxially stretched and also in only one lateral direction

called uniaxially stretched. For uniaxially stretched PVDF film, the lateral strain constant

in the direction of stretch is high (i.e. d31 >> d32). For the biaxially stretched film, the

lateral strain constant is almost equal (i.e. d31 = ~d32). The stretching changes the

mechanical properties and density of the material. The dielectric and piezoelectric

properties of PVDF are temperature dependant. While the electromechanical coupling

63

coefficient kt is independent of temperature, the dielectric properties and mechanical loss

tangent are temperature dependant. When used in imaging transducer, the center

frequency and bandwidth vary with temperature. The acoustic velocity of PVDF

decreases approximately linearly with temperature. So the acoustic impedance and the

resonance frequency of the imaging transducer made with PVDF decrease similarly with

temperature. PVDF is highly flexible and thus allows application on curved structures. It

can achieve large strain because of its high depolarization field and high field strength

against electrical breakdown.

One of the polymers used in transducer application is polyvinylidene fluoride (PVDF),

which is semicrystalline. PVDF is commercially available as thin films ranging from 4 –

110 µm thick except in special cases where much thinner/thicker films are made. Owing

to poling difficulties, the thickest available film is usually 110 µm except in special cases

where higher thickness films are made. The advantages of this material are that it is

wideband, flexible, and inexpensive. Generally, the use of matching layers limits the

bandwidth. Because the acoustic impedance of PVDF (4 MRayls) is close to that of

water (1.5 MRayl) and muscle (1.7 MRayl), the reflection at the interface between

imaging transducer and the medium of propagation is minimized. Hence matching layers

are not necessary, thus preserving the intrinsic broadband property (mechanical quality

factor Qm = 1/tanδm) of PVDF. The property of large mechanical internal loss

(mechanical loss factor tanδm = 0.10) of PVDF decreases the response time considerably

and improves the sharpness of the response, thus contributing to widening the frequency

bandwidth of the polymer transducers [45]. The short impulse response leads to high

64

spatial resolution capabilities in medical imaging. Their low radial/lateral mode coupling

reduces the effects of edge waves and near field distortions, which is typical of piezo

ceramic transducers. Research has shown that planar PVDF transducers can produce

plane wave performance in the acoustic near field, which is superior to that of

piezoceramics [44].

The disadvantages are (i) that it has a low thickness mode electromechanical coupling

factor (kt = 0.20), which governs the transmitting efficiency. The square of this number is

a measure of the ratio of the transformed electromechanical energy to the total input

energy. Therefore, with equal amount of available input electrical power, the PVDF

transducer will generate less acoustic output compared to PZT. Larger kt contributes to

the increase in the response amplitude to a great extent. (ii) It has large dielectric loss

factor (tanδe = 0.25), which will result in an appreciable amount of electric power being

dissipated in the transducer itself. (iii) It has low relative dielectric constant (6), which

will result in very high input electrical impedance in comparison to that of a PZT.

Therefore, to generate equal amount of output acoustic power, a higher drive voltage is

required for PVDF transducer. The capacitance of the transducer is proportional to the

dielectric constant, which implies that the voltage developed across the transducer in the

receiving mode is inversely proportional to its dielectric constant. Low dielectric

constant is advantageous in terms of developing large voltage signal.

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3.4 Modes of Vibration

Although the performance of the imaging transducer depends mostly on the kind of

piezoelectric material, other factors such as frequency, backing material, matching layer

and the geometry of the active element are also have their influences on the efficiency of

the transducer. The acoustic energy from the piezoelectric material is due to the

vibrations of material. The mode of vibration is related to the geometry of the

piezoelectric material. Normally the vibration is required in one direction only. The

vibration in other directions if any will cause artifacts and energy loss.

Usually, there are three modes of vibration namely (i) plate mode (ii) thickness mode and

(iii) bar mode. The schematic illustrations of three important vibrational modes are

shown in Figure 3.

w

tt

l

l

tw w

l

Figure 3: Geometry of three important vibration modes

Plate mode w, l >> t

Thickness model >> w

Bar mode t >> l, w

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Plate mode

In the plate mode, the width (w) and length (l) are much larger than the thickness (t) of

the piezoelectric material. If the vibration is in the width direction, then it is called lateral

mode.

Thickness mode

In the thickness mode, the length (l) is much larger than the width (w) and thickness (t).

For optimum efficiency, the ratio of width (w) and thickness (t) should be smaller than

0.7, if the thickness vibrational mode alone is required.

Bar mode

In the bar mode, the thickness (t) is much larger than the length (l) and width (w).

In all the vibrational modes, the poling is done in the thickness direction and the

metalization is normal to the poling direction. The velocity in the piezoelectric material

is related to its elastic properties. The acoustic velocity is different for different modes of

vibration. The piezoelectric ceramic used should satisfy certain relationship between

length, width and thickness in order to have only one of the vibrational modes. If the

prescribed relationship is not maintained, then there will be more than one vibrational

mode, which results in poor performance. All imaging transducers are made to operate in

thickness mode.

3.5 Summary

In order to achieve broad bandwidth and high sensitivity simultaneously, the following

properties of the piezoelectric materials should be considered: (a) coupling factor (b)

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dielectric constant (c) losses (d) depoling temperature and (e) velocity. The depoling

occurs whenever the piezoelectric material approaches the Curie temperature. The

piezoelectric materials with high velocity give high impedance. The various parameters

of the piezoceramic and piezopolymer materials are compiled from various publications

and given in Table 1 for easy comparison.

Table 1. Comparison of piezoelectric material parameters SL No.

Material Property PiezoCeramic (PZT5H)

PiezoPolymer (PVDF)

1 Transmitting Constant: d (10-12 C/N) K583 K15 2 Receiving Constant: g (10-2 Vm/N) K1.91 K14 3 Electromechanical coupling coefficient: kt

K0.55 F(0.15…0.2) 4 Free dielectric constant: εT (10-11 F/m) K3010 K9.7 5 Clamped dielectric constant: εS (10-11 F/m) L531 B5 6 Acoustic velocity: c (m/sec) K3970 B2200…O2260)7 Density ρ (kg/m3) K7450 O1780 8 Acoustic Impedance (Z) kg/m2-s*106 P33.7 O4.02 9 Electrical loss tangent: tanδe L0.002 O0.25 10 Mechanical loss tangent: tanδm L0.004 O0.1 11 Curie Temperature 0C. K190 K100 12 Typical –6dB fractional bandwidth K60% D>150%

Note: These data were compiled by the author from various sources of literature: KFrom[9], OFrom [45], BFrom [44], FFrom [32], DFrom [46], LFrom [47 ]

To compensate for the high electrical impedance of the piezopolymer, use of multilayer

structure, passive tuning, placing an impedance transformer or active preamplifier near

the element are the possible options.

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3.6 Conclusion

By studying the properties of the several types of piezoelectric material, it was decided to

explore the feasibility of using piezopolymer for the imaging transducer because of the

following reason. The piezoceramic provides good pulse echo sensitivity but relatively

narrow bandwidth. Single crystal demonstrates best sensitivity and wide bandwidth

compared to piezoceramic. But the piezopolymer provides the highest bandwidth but

poor sensitivity. By employing multilayer structure and utilizing the concept of Barker

code, high sensitivity and highest bandwidth can be simultaneously achieved using

piezopolymer.

The next chapter deals with the development of double layer polymer hydrophone probe,

which not only aimed in achieving simultaneous improvement of bandwidth and

sensitivity but also provided experience in fabricating complex multilayer Barker code

transducer as described in chapter VI.

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CHAPTER 4: DEVELOPMENT AND CHARACTERIZATION OF IMPROVED DESIGN OF POLYMER HYDROPHONE PROBE

A detailed description of the development process, performance evaluation, results and

discussion about the newly developed double layer polymer hydrophone probe is

presented in Appendix A. However for the purpose of continuity for the readers and sake

of completeness, a brief description of the characteristics of the double layer hydrophone

probe is presented in this chapter.

4.1 Introduction

As a part of research oriented product development work, several wideband double-layer

ultrasonic polymer hydrophone probes with different thicknesses of PVDF film having

different active diameters were fabricated and tested for several of its performance

characteristics such as acoustic sensitivity, bandwidth, frequency response, angular

response, effective aperture size and orientation effects at Perceptron, Inc. in order to

explore the fundamental improvement in the existing design [48, 49]. The hydrophone

probes were tested at the in house laboratory. The development of this polymer-based

hydrophone is specifically intended in evaluating the diagnostic ultrasound imaging

transducers.

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4.2 Synopsis

4.2.1 Acoustic sensitivity

The sensitivity of a double layer hydrophone is determined by the resonance of the

membrane film, and the electric and piezoelectric properties of the PVDF film. The

sensitivity of a particular hydrophone is determined by its active area and the capacitive

loading of the element, electrode leg and connecting cable. The evaluation of the

performance characteristics of several hydrophones revealed that larger spot size

hydrophones exhibited better sensitivity.

4.2.2 Frequency response and bandwidth

The double-layer hydrophone probes made by employing 25 µm and 9 µm thick PVDF

films having 0.4 and 0.6 mm of geometrical spot diameter were tested for their

performance. The results of the randomly chosen hydrophone probes were also verified

at National Physical Laboratory (NPL), UK. The frequency response of a double-layer

hydrophone probe fabricated with (9+9) µm thick PVDF film measured at NPL showed

that the sensitivity is being constant to ±3 dB between 1 to 50 MHz. In the case of

hydrophone probe fabricated with 25+25 µm thick PVDF film, the sensitivity variation in

the frequency range from 1 to 20 MHz is within ±3 dB of the mean value for all

frequencies. The double layer polymer hydrophones developed using dissimilar

thickness (25+9 µm) were demonstrated higher bandwidth compared to that of

hydrophones made using similar thickness (25+25 µm) of PVDF film. A flat frequency

response up to 25 MHz has been demonstrated.

71

4.2.3 Angular response

The angular responses of several double-layer PVDF hydrophones of different thickness

of PVDF film and various geometrical spot sizes were obtained by rotating the

hydrophone in the far field of the transducer’s plane wave and measuring the

hydrophone’s response at some angle of rotation. The angular response and the

frequency dependency were studied.

4.2.4 Effective aperture size

The effective diameter of the hydrophone was determined for different thickness and

diameter hydrophones at varies frequencies as the determination of effective diameter is

necessary for applying the spatial averaging corrections [50] while reporting the acoustic

measurement results of medical ultrasonic transducers.

4.2.5 Orientation effects

To study the apodized behavior and asymmetry of the sensitive region of the hydrophone

probes, angular response of several hydrophone probes of different kinds, various

geometric spot sizes and different thicknesses of PVDF film were measured in two

rotational axes perpendicular to each other. The effective diameter of the hydrophone

probe was determined from the measurement of its directivity pattern. From the results,

it was seen that the effective diameter of the hydrophone varied significantly with

reference to axis of rotation. There was a marked asymmetry in the effective diameter

measured along the two orthogonal axes. The cause and effect of such asymmetry were

investigated.

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4.3 Conclusion

This improved design of double-layer hydrophone probes to enhance the bandwidth using

dissimilar thicknesses of (25+9 µm) PVDF film developed in this work was the first one

of this kind and no one had developed this type previously. The new design hydrophone

exhibited higher bandwidth compared to 25+25 µm thick film and possessed better

sensitivity and robustness compared to 9+9 µm thick PVDF film. This experience helped

the furtherance of understanding of the behavior of polymer probes, as explained in the

previous sections. The drop method of adhesion procedure used in fabricating the

hydrophone was adopted in developing multilayer imaging transducers after optimizing

the glue thickness.

The next chapter deals with the new measurement technique developed for calibrating the

double layer hydrophone below 1 MHz, which is essential in accurately characterizing

the newly developed imaging probes before use in clinical practice.

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CHAPTER 5: DEVELOPMENT OF CALIBRATION PROCEDURE

A detailed description of the development of a new measurement technique along with

the results and discussion are presented here. Development of a new measurement

technique is essential in view of the non-existence of such a procedure in calibrating the

hydrophone below 1 MHz, which is important in characterizing the newly developed

imaging transducers for assessing the potential bioeffects while in clinical use.

5.1 Introduction

The primary motivation of this work was to develop and optimize a rapid and reliable

ultrasonic hydrophone calibration procedure suitable for use in the frequency range from

100 kHz–1000 kHz and to investigate the feasibility of using a planar scanning technique

to determine the hydrophone’s sensitivity in this frequency range. The outcome will

facilitate establishing the foundation for the implementation of low frequency calibration

technique for diagnostic and therapeutic ultrasound equipment, including High Intensity

Focused Ultrasound (HIFU) transducers.

5.2 Calibration Setup

Two piezoelectric (PZT) transducers (Panametrics and KB-Aerotech) were used as

acoustic sources; Panametrics’ source operated at the center frequency of 0.5 MHz and

the KB-Aerotech’s at 1 MHz. The results were obtained in the following way. The

source transducer was placed in a tank containing deionized and degassed water at room

temperature with the transducer face oriented downwards. The use of deionized water is

advisable to minimize the effect of possible corrosive interaction between the water and

74

hydrophone housing and metal electrodes. IEC standard recommends that the water

conductivity should be less than 10 µSiemens/cm [23]. The degassed water was used to

minimize formation of micro bubbles on the hydrophone and source surfaces, which not

only perturb the ultrasound field during measurements but also lead to cavitational effect.

An automatic XYZ scanning system, controlling the hydrophone and the source

transducer movements through stepping motors was used. The hydrophone under test

was supported by a holder featuring independent translations in the X, Y and Z

directions. The source was also mounted on a XYZ manipulator and could be tilted and

adjusted about the vertical Z-axis. The oscilloscope and stepper motor controller were

interfaced to a personal computer through an IEEE-488 bus for capture and storage of the

hydrophone signals [51].

During the calibration, the source transducer was excited with a 10 cycles tone burst

generated by a HP 8116A signal generator and amplified by a linear wideband power

amplifier (ENI 350L). Tone burst transmission was used to avoid the undesirable effects

of standing waves and multiple reflections from the hydrophone and the walls of the tank.

Nonlinear propagation effects were avoided by keeping the transmitter driving voltage

sufficiently low so that the second harmonic of the received signal was at least 30dB

below the fundamental (center) frequency. The 0.5 MHz source transducer was driven at

0.3, 0.35, 0.4, 0.5, 0.6 and 0.65 MHz whereas the 1 MHz transducer was driven at 0.65,

0.7, 0.75, 0.8, 0.9 and 1 MHz to cover the whole frequency range considered. Thus at the

frequency of 650 kHz, it was possible to obtain data produced independently by the two

sources. As evidenced from the results, the 650 kHz data are in good agreement (to

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within ± 0.7 dB), which further enhances confidence in the robust nature of the planar

scanning technique.

5.3 Calibration Procedure

5.3.1 Initial alignment procedure

One of the double layer hydrophone probes fabricated by me, which has a preamplifier,

was calibrated at the National Physical Laboratory (NPL), UK at 0.25, 0.5, 0.75 and 1

MHz and was used as a reference hydrophone. The working hydrophone used in the

measurements described below was another double layer hydrophone but without

preamplifier. The working hydrophone was cross calibrated against the above reference

hydrophone at discrete frequencies by substitution. The calibration by substitution was

performed in the far field of the source transducers.

The hydrophone was placed in the water tank approximately 30 minutes before the

measurements began. The position of the source transducer was adjusted to maximize

the output signal from the hydrophone, which was displayed on the Tektronix 2440

digital oscilloscope. Once the actual acoustic axes of the source and the hydrophone

were aligned, the pressure-time waveform corresponding to the maximum output signal

was captured and Fourier transformed to determine the center frequency (arithmetic mean

of the upper and lower half power frequencies) [26]. The center frequency was used to

determine the working hydrophone sensitivity, corrected for the oscilloscope input

impedance (1 MΩ 15 pF).

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5.3.2 Determination of the peak pressure amplitude location

Next, the working hydrophone was scanned axially at 1 mm increments along the

acoustic axis in the far field region. The received signal was displayed on digital

oscilloscope (Tektronix 2440) to detect the peak voltage and the influence of noise

present in the measurement system was minimized by using the signal averaging facility

of the oscilloscope. Once the axial location that produced the highest in-water intensity

was determined by computer analysis, the working hydrophone was repositioned in the

location corresponding to the far field maximum pressure amplitude. The working

hydrophone signal was re-maximized and the pressure-time waveform was captured.

Based on this waveform, the intensity values and other relevant parameters were

calculated. Before each measurement, the peak-to-peak driving voltage was measured at

the source terminal to ensure identical excitation conditions.

5.3.3 Calculation of pulse intensity integral

The Pulse Intensity Integral (PII) (see Fig 1b) was calculated using the following

expression [14, 45]:

(5.1)

where p (t) is the instantaneous acoustic pressure, ρ is the density, c is the speed of sound

and the integration limits bound the pulse. The value of the PII is a measure of the total

energy in the pulse. The pulse duration was determined by finding the time for which the

PII is between 10 and 90 percent of its final value. This time was multiplied by 1.25 as

called for in the NEMA Standard [12]. Knowing the PII, the pulse duration, and the

dttp∫= c

)(PII2

ρ

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pulse repetition frequency, the pulse average and time average intensity were calculated

[12]. An example of the representative plots showing the hydrophone received voltage

waveform and corresponding pulse intensity integral together with the frequency

spectrum of the signal received from the working hydrophone are shown in Figure 4.

These plots correspond to a tone burst excitation at 1 MHz. A similar set of plots was

obtained for each frequency considered here.

Figure 4: Typical measured hydrophone waveform with the corresponding PII and frequency spectrum of 1 MHz circular piston source.

78

5.3.4 Cross axis scan

Cross axis scans (lateral beam profiles) shown in Figure 5 were conducted to determine

beamwidth of the acoustic sources so that appropriate sampling interval and number of

sampling points sufficient to measure the beam down to the –26 dB level could be

identified while performing the raster scan [51]. The procedure involved spatial

integration of the beam intensity values in the acoustic field along two perpendicular

directions. The profiles of Figures 5a and 5b were obtained in the following way. The

hydrophone was moved in a plane perpendicular to the acoustic axis of the source from

the spatial peak location to the -26 dB point location. Next, the hydrophone was scanned

back through the maximum of the beam to the -26 dB point on the other side of the beam.

The procedure was repeated to obtain a beam width in the plane perpendicular to the one

already scanned. The signals were averaged to improve the signal-to-noise ratio, and the

pressure-time waveform was captured at each measurement point, transferred to the

computer and analyzed. The plots shown in Figure 5 indicate that the measured –6 dB

beamwidths (solid line) are almost identical with the theoretically predicted Gaussian

shaped ones (dashed line). However, due to the finite aperture of the active transducer

element and the influence of the side lobes in practice, below the –6 dB level, the

Gaussian shape of the overall radiation pattern is not retained.

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5.3.5 Raster scan

The raster scan was performed to calculate the total acoustic power produced by the

source. The scan plane was scanned in a two-dimensional raster pattern, parallel to the

sinusoidally excited plane piston source transducer over an array size of 30x30 with a

sampling interval of 1.5 mm. This sampling interval (step size) was chosen based on the

results of the cross axis scan measurements as described above. The distance between

the sampling point is usually less than or equal to one wavelength or the hydrophone

diameter, whichever is higher. The sampling was performed down to the –26 dB level

for the maximum measured in the given plane to ensure that the active cross section of

Figure 5: Beam plot along the x and y axes of the 1 MHz acoustic source. The dashed lines show the gaussian beam (theoretical) distribution with the same –6 dB beamwidth as that of the measured transducer.

80

the source could be identified [51]. At each sampling point, the hydrophone output

voltage was captured, peak detected and transferred to the computer memory.

Concurrently the corresponding pulse intensity integral (PII) was computed. The

resulting data set was stored as a two-dimensional matrix representing the spatial pressure

variation over the scan plane. The surface integral was then obtained by rectangular

numerical integration [30]. The experimental setup for the raster scanning and the

illustration of the raster fashion are shown in Figure 6.

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Figure 6: Experimental set up for raster scanning (A) and illustration of raster fashion (B)

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Raster scan beam intensity distribution for the 1 MHz sources is shown in Figure 7.

The set of measurements described above and needed to obtain the total acoustic power

produced by the two sources used was made at 11 frequencies (i.e. 300, 350, 400, 500,

600, 650, 700, 750, 800, 900 and 1000 kHz). This number of measurement points was

considered appropriate to obtain an adequate sensitivity versus frequency response in the

range 300 – 1000 kHz. As already noted, since a single wideband source in the

frequency range 0.1 MHz to 1 MHz was not commercially available, two transducers

having nominal center frequencies of 0.5 and 1 MHz were used as acoustic sources.

Rel

ativ

e In

tens

ity in

dB

0

-6

-12

-24

-30

-18

Figure 7: Three-dimensional representation of the intensity field produced by the 1 MHz source transducer and obtained from planar scanning

Y-axis

X-axis 30x30 Samples at 1.5 mm spacing

83

5.4 Radiation Force Balance Measurements The Radiation Force Balance (Cahn C-32) measured the radiation force of the beam

striking a totally absorbing target (SOAB and HFMA). The diameter of the target was

larger than that of the source transducer and in this case, the total beam power could be

determined directly by measuring the force exerted on the target. The radiation force

pressure F, was determined as:

F = cθcosP (5.2)

where θ is the angle of incidence of acoustic beam, P is the magnitude of the time

averaged, spatially integrated power intercepted by the target and c is the speed of sound

in water. With θ = 00 (normal incidence) and for an absorbing target, F/P= 67 mg/W. i.e.

one watt of ultrasound power induces a force equivalent to that of 67-milligram weight

[52]. In measurement practice, a settling time of an hour after filling the balance with

degassed water and inserting the transducer is usually recommended to ensure that any

possible convection currents have been minimized and air bubbles, if any, on the target or

transducer have been reabsorbed into the water. The power measurements were repeated

5 times to take into account thermal drifts and minimizes uncertainty of the

measurements. The next section describes calibration of the selected hydrophone in

terms of µV/Pa based on the measurements of total acoustic power and raster scan beam

distribution.

84

5.5 Hydrophone Sensitivity Calculation

The basic procedure underlying the planar scanning method of calibration involves in

producing a known pressure field at the hydrophone and measuring the hydrophone’s

electrical output. The core of this procedure is to assess the pressure level of a known

acoustic power of the source transducer independently by using the hydrophone. The

“intensity response factor” is one form of expression for the sensitivity of the

hydrophone, which is the relationship between the plane wave intensity at the

hydrophone and the square of the voltage produced at the hydrophone terminal as a result

of that intensity. Since the planar scanning method and the radiation force balance both

measure the acoustic output of the source, by comparing the power obtained from the

raster scan measurements with the power measured with the Radiation Force Balance

(RFB), the intensity response factor ( 2fK ) of the hydrophone was calculated as [30]:

RFBPower

dttVPRFK

T

TP

f

∫= 0

2

2

)(* * ∑∑

= =

∆∆ M

i

N

jjiTP

TP

yxVV

yx1 1

22 ),(

)0,0( (5.3)

where

M = Number of linear scans

N = Number of sample points per linear scan

x∆ = Distance between linear scans

y∆ = Distance between sample points

VTP = Temporal peak voltage

PRF = Pulse Repetition Frequency

85

Equation (4) indicates that the units of Kf2 are V2 W-1 cm2. Since hydrophone probe

sensitivity Mh is usually given in terms of V/Pa or dB re 1µV/Pa, the values of Kf2 were

converted into the corresponding values of Mh using the following relationship [53].

Kf2 = 1.5 * 1022 * [10(M (dB)/10)]………( V2 W-1 cm2) (5.4)

Mh = 10 log (10 * 1.5 22

2fK

) ………….. (dB re 1V/µPa) (5.5)

It should be noted that this relationship holds for plane waves. To comply with this

requirement, the measurements were carried out in the far field of each source. The

results obtained are summarized in the next section.

5.6 Comparative Study

To facilitate a comparison and discussion of the calibration results obtained using the

planar scanning technique, calibration of the same hydrophone using Time Delay

Spectrometry (TDS) technique combined with substitution calibration method [2-4] was

performed. The results were also verified with the data provided by National Physical

Laboratory (NPL), UK, using independent calibration.

86

5.7 Results Correlation and Discussion

The plots of Figure 8 summarize the end of cable (EOC) voltage sensitivity versus

frequency for the working membrane hydrophone. The calibration results obtained using

the planar scanning technique are presented together with those provided by the

independent national laboratory (NPL, UK) and also with those determined by

substitution using the swept frequency technique [28, 54]. Since the swept frequency

calibration technique yields virtually continuous frequency response, for clarity and

consistency the error bars corresponding to discrete planar scanning technique data and

NPL data are not shown in Figure 8.

-270.0

-269.0

-268.0

-267.0

-266.0

-265.0

-264.0

-263.0

-262.0

-261.0

-260.0

0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1

Frequency (MHz)

Volta

ge S

ensi

tivity

(dB

re 1

V/uP

a)

Planar Scanning data

NPL Calibration data

Swept Frequency Calibration data

Figure 8: End-of-cable voltage sensitivity versus frequency of the double layer PVDF hydrophone in The frequency range 0.3 to 1 MHz obtained using Planar scanning technique (), Time delay spectrometry method (—) and calibration data provided by National Physical Laboratory (◊).

87

Based on the results presented in the previous section, the experimental data indicated

that the sensitivity variation of the double layer hydrophone probe by employing planar

scanning technique was approximately ±1 dB from 0.3 to 1 MHz [5]. The calibration

results of the working hydrophone by using the TDS technique combined with

substitution method [3, 4] agreed well with those obtained here using planar scanning

approach. The overall uncertainty of the TDS calibration was estimated to be better than

±1.5 dB in the frequency range considered. The data provided by NPL at discrete

frequencies using independent calibration method exhibited a sensitivity variation of

approximately ±1 dB, which agrees very well with the calibration results obtained by

employing planar scanning technique and TDS technique in this work. Such uncertainty

is considered to be acceptable in the underwater acoustic measurements carried out in a

similar frequency range [55]. Hence the planar scanning approach developed in this

work offers an alternative to other primary calibration techniques such as reciprocity or

broadband pulse technique. Planar scanning technique provides discrete frequency

calibration and it would be desirable to produce additional data in the frequency range

overlapping the upper bandwidth of the 0.5 MHz source and the lower bandwidth of the

1.0 MHz source. However, the transmitting voltage sensitivities of both sources were

insufficient to provide adequate signal to noise ratio. On the other hand, as already noted

in section 2.2, the data obtained at 650 kHz were in good (to within ±0.7 dB) agreement,

which further enhances confidence in the robust nature of the planar scanning technique.

In general, it would be desirable to use a single wideband transducer source capable of

producing a signal in the whole frequency range considered. The availability of such

88

transducer would facilitate implementation of rapid hydrophone calibration. While such

wideband source, producing a signal with adequate signal to noise ratio, is not

commercially available, it was built and described in [54], where the TDS technique was

used to provide a near continuous frequency response of both membrane and needle

hydrophone probes in the frequency range 0.25 to 2.5 MHz [55]. In this work it was

desirable to obtain calibration data down to 100 kHz, however, the calibration could not

be performed below 300 kHz due to the limitation of the available equipment (Power

Amplifier ENI 350L) and inadequate signal to noise ratio achieved with a low frequency

source operating at frequencies below 0.3 MHz. The results obtained were also

examined to determine the possible spatial averaging effects. As the effective diameter

of the hydrophone increases with decreasing frequency, the spatial averaging correction

was not required in the measurements reported here because the ultrasound beam was

sufficiently wide to cover the active element of the working hydrophone in the frequency

range considered.

5.8 Summary

To address the issue of non-existence of calibration standard and the need for a new

calibration procedure, a planar scanning technique was developed to determine the

frequency response of ultrasonic polymer hydrophone probes below 1 MHz and was

compared with the sensitivity data determined by TDS calibration technique combined

with the substitution method [31] and the data obtained by using independent calibration

method at National Physical Laboratory (NPL), UK, which is an internationally

recognized measurement laboratory. The experimental data indicate that the sensitivity

89

variation of the hydrophone probe determined by employing the planar scanning

technique [5] is approximately ±1 dB, which agrees well with the results obtained by

using Time Delay Spectrometry [TDS] technique [28, 54] and with the data provided by

NPL. The double layer hydrophone studied had bandwidth extending below 0.25 MHz.

The results are shown in Figure 8. Harris and Gammell measured the voltage sensitivity

of piezoelectric hydrophones in the frequency range from 0.2 – 2 MHz using a broadband

pulse technique and reported that the sensitivity of the bilaminar membrane hydrophone

[27] was essentially flat. Their results were supported by the results obtained in this work

by using planar scanning technique [5] and TDS technique [28, 54]. The planar scanning

technique is suitable for absolute calibration of hydrophone probes in the frequency range

below 1 MHz to within ±1 dB and offers an alternative to other primary calibration

technique such as reciprocity or broadband pulse technique [5]. So far, no one has

explored this planar scanning technique for calibrating the hydrophones below 1 MHz.

5.9 Conclusions

In conclusion, a calibration method using planar scanning technique was developed to

determine the low frequency response of the PVDF ultrasonic hydrophone probes. This

method is suitable for absolute calibration of hydrophone probes in the frequency range

below 1 MHz to within ±1 dB and offers an alternative to other primary calibration

techniques such as reciprocity [5, 55] or broadband pulse technique [27]. The method

developed could also be used while characterizing the newly developed ultrasonic

transducers for medical imaging in accurate measurement of peak rarefactional pressure

at frequencies down to fc/20 as required by federal regulation and thus in arriving precise

90

determination of Mechanical Index (MI). It would appear that the method developed

could also be used in characterization of medical and therapeutic ultrasound devices

operating at frequencies below 1 MHz and for periodical spot check of working and

reference hydrophones.

The next chapter addresses all the aspects of the development and testing of single

element, single and multilayer imaging transducers and one of the research goals of

simultaneous improvement in sensitivity and bandwidth is dealt with.

91

CHAPTER 6: DEVELOPMENT AND CHARACTERIZATION OF SINGLE ELEMENT NON-RESONANT POLYMER IMAGING

TRANSDUCER

The procedure devised and process developed for fabrication, the related design and

development issues, and the results and analysis of the polymer imaging transducers are

presented in this chapter in detail starting from the material characterization.

6.1 Introduction

In medical ultrasound imaging, fundamental improvements should result in both the

improved image resolution and increased penetration depth. This depends primarily on

simultaneously achieving broader bandwidth and higher sensitivity in a single transducer.

In order to achieve this goal of simultaneous improvement of bandwidth and sensitivity,

piezopolymer was explored as a transducer material for imaging transducer. The

multilayer structure with Barker code approach was implemented in fabricating the

transducer and the results of multilayer structure were compared with a single layer

transducer.

6.2 Characterization of PVDF film - Impedance/Admittance Measurements

To validate the resonance and non-resonant behavior of the piezo film and PVDF

transducers respectively, the fundamental thickness mode frequency measurements were

performed using an Agilent 4395A Network Analyzer. The electrical input admittance

and impedance measurements were carried out separately (for the sake of counter

verification) on an electroded PVDF film of 56 µm thick and having 3/8” diameter. The

half wavelength fundamental thickness mode resonance frequency of PVDF film was

92

observed to be 19.8 MHz while the theoretical prediction was (2200 m.sec-1/(2*56 µm))

19.6 MHz. The resonant behavior of the PVDF film is clearly seen from the admittance

and impedance plot shown in Figure 9 and 10.

6.3 Design and Development of System Electronics and Controls

A common problem with all pulse-echo imaging systems is the need to couple the

transducer to both the high voltage excitation pulse and the sensitive receiving

electronics. A transmit-receive (T/R) switching circuit by using two pairs of crossed

diodes was built, in which the diode pairs were closed during the application of higher

voltage (> 0.5 V) pulse but open while the probe was receiving energy. During the

excitation pulse, a clear signal path was established between the excitation source and the

transducer, while the input to the preamplifier is shorted to ground. After the excitation

pulse has passed, both pairs of crossed diodes become non-conducting. A high frequency

0.00

0.01

0.02

0.03

0.04

0 10000000 20000000 30000000Frequency in Hz

Adm

ittan

ce m

agni

tude

in M

hos

0

20

40

60

80

Phas

e in

Deg

ree

fr=19.8MHz

Admittance magnitude plot

Admittance Phase plot

Figure 9: Plot showing the variation of admittance magnitude and phase of a 56 µm piezo film in air

Figure 10: Plot showing the variation of impedance magnitude and phase of a 56 µm piezo film in air.

0

400

800

1200

1600

0 10 20 30Frequency in MHz

Impe

danc

e M

agni

tude

in O

hms

-80

-60

-40

-20

0

Impe

danc

e Ph

ase

in D

egre

es

Impedance magnitude plot

Impedance phase plot

fr = 19.8 MHz

93

(200 MHz) National Semiconductor current feedback amplifier (CLC 411) was selected

for the preamplifier and the circuit built and tested for satisfactory operation.

6.4 Design Considerations of Single Element Polymer Imaging Transducer

6.4.1 Material selection, thickness and active area

PVDF was selected as the basis for a wideband transducer due to the relatively high

operating frequency, wide potential bandwidth, reasonable acoustic impedance matching

to water and tissue, mechanical flexibility and robustness. The thickness and area of an

active piezoelectric polymer resonator are important criteria for transducer operating

frequency, bandwidth and insertion loss. The operating frequency of the transducer is

determined by the thickness of the film and the boundary condition. The thickness is

usually chosen as λ/2 and it should be at least an order of magnitude smaller than the

lateral dimension. A half wavelength resonator design was selected to maximize the

bandwidth and it was achieved by using a low impedance backing material. The

thickness of the PVDF material was selected as 56 µm for the desired frequency range of

2-15 MHz. Circular PVDF transducers were proposed for the single element transducer

in order to have minimal edge diffraction and hence may be considered as an acoustic

source producing a plane wave thereby minimizing possible response variation due to

aperture diffraction [57]. Although a large active area will have a better electrical

impedance match between the transducer and the driving electronics, smaller aperture

with large depth of field is generally preferred in medical ultrasound imaging. But the

small aperture will lead to higher impedance resulting in poor sensitivity of the probe.

94

6.4.2 Electroding

The single element imaging transducer was designed with a circular electrode having a

diameter of 3/8” (using a thin 56 µm film) of Ni+Cu+Ni (1000 Angstrom thick), which

provide adequate performance without acoustically loading the piezoelectric polymer

much. The film was electroded on both sides. The active area of the film is determined

by the overlapping electrodes. The electrode patterns on opposing surface of the films

are defined in such a way that by bonding the film, a multilayer structure is created. Due

to mass loading effects, thickness layer of electrodes has some loading and damping on

PVDF and as a result one expects lowering of the resonant frequency of the transducer.

6.4.3 Backing material

Selection of the proper backing material for a piezoelectric polymer ultrasonic transducer

demands careful consideration of such parameters as operating frequency, bandwidth,

insertion loss, and operational environment. While using brass as backing material for

PVDF which has an impedance of ~32 MRayals, acoustic reflections at the brass-PVDF

interface have 1800 phase shift relative to the reflections occurring in probe using ceramic

as piezoelectric material. Therefore, PVDF transducer with brass backing resonates in a

λ/4 thickness mode. When PVDF is backed by a material having an acoustic impedance

equal to or lower than the PVDF, the transducer will resonate in λ/2 mode, just like

ceramic probes. In order to achieve a broadband operation, the Q factor must be low.

The Q factor depends on the impedance of the backing and propagation medium. By

choosing a lossy backing material of appropriate impedance, the Q factor can be lowered.

In general, the acoustic requirements for an absorber backing are (i) it must have an

95

acoustic impedance that is close to that of the piezoelectric polymer and (ii) it must have

sufficient acoustic attenuation to prevent unwanted acoustic reverberations (i.e. back wall

reflections). Other important requirements for the backing material include, that it (i)

must be able to adhere to the piezoelectric material and (ii) must be available in

sufficiently thick substrate form and have a high surface quality (e.g. polished). Since the

acoustic impedance of PVDF material is about 4 MRayls, backing for the piezoelectric

film is provided by using a non-piezoelectric rod made from Kynar polyvinylidene

fluoride (PVDF) resin having a measured acoustic impedance of 3.81 MRayls. The low

impedance backing moves the resonance frequency to a higher value near towards the

value for half wavelength film thickness, compared to high impedance backing. The

backing block was carefully shaped and polished on the front surface, whereas it was

roughened on the back surface to minimize internal reverberation.

6.4.4 Wear protecting front matching layer

The imaging transducers generally require a front layer for wear protection. For the

transducer developed in this work, which is operating at half wave resonance frequency,

the thickness of the front layer should be λ/4 at its resonance frequency, which is

estimated as:

tFL = tPVDF * PVDF

FL

cc

2 (6.1)

where, tFL and cFL are the required thickness and the longitudinal speed of sound in the

front layer material and tPVDF and cPVDF are the thickness and speed of sound in the PVDF

film respectively. The proper front layer material will improve the sensitivity but

decrease the bandwidth of the transducer slightly. Since the acoustic impedance of the

96

PVDF and tissue are about 4 MRayls and 1.6 MRayls respectively, the optimum front

layer material of the imaging transducer should have an impedance of about 2.5 MRayls

( 6.1*4 ). The front layer can be designed in such a way that it will serve both for

impedance matching and as well as focusing. To avoid quarter wavelength filtering

effects, the front surface was not covered with any protective coating to the transducers

developed in this work.

6.4.5 Adhesion

The adhesive used for gluing the PVDF film and the backing must have acoustic

impedance close to that of PVDF. This will prevent reflection at the boundary of the

acoustically mismatched material. The adhesive layer must be extremely thin (very much

less than the wavelength = ~1/10 of wavelength) such that the capacitance of the adhesive

layer is many times greater than that of the active piezoelectric layer. Otherwise, the

adhesive layer introduces an undesirable series capacitance voltage divider and

undesirable acoustic reverberations and losses. The thickness of the glue will have an

impact on the resonance frequency, as the resonance frequency is determined by:

)(*22/

adhesivefilm ttcf+

=λ (6.2)

where c is the speed of sound in the PVDF material.

6.5 Fabrication Process

The knowledge gained and the technique of lamination adopted in the fabrication of

double layer ultrasonic hydrophone probes were applied in developing the multilayer

polymer acoustic transducers. Three layer polymer transducers were fabricated using 56

97

µm thick PVDF film in such a way that it produces Barker coded pulses when excited.

This was achieved by arranging the polarization pattern of the PVDF films according to

the Barker code length 3. The fabrication of Barker code multilayer transducers

facilitated in validating the Barker code concept experimentally. A single layer

transducer was developed to demonstrate the wideband characteristics and compared with

a multilayer transducer. The brief description of the fabrication process is given below.

A uniaxially stretched and poled PVDF sheet of 56 µm thickness and having circular

electrode of 10 mm diameter, which had been sputtered on to the PVDF film with

Ni+Cu+Ni (1000 Angstrom thick) was acquired from Measurement Specialities, Inc.,

Norristown, PA.

The electrical contacts were established on the electrodes by bonding a thin wire with a

conductive silver epoxy. A non-piezoelectric cylindrical rod of polymer (Kynar) was

polished smoothly on the gluing side and the other side was roughened to prevent

acoustic reverberation. Polymer backing was chosen in order to obtain fundamental

thickness resonance of half wavelength. After experimenting with several brands of

adhesives, a non-viscous epoxy RBC 3200 from RBC Industries, Inc. was chosen for

gluing the film, which has a low viscosity of 300 cps. Manual applying of epoxy on one

side of each film was attempted, but it did not yield the desired thin and uniform

bondline.

98

In order to produce a uniform and thin bonding layer, a bonding technique called “drop

method” that was used for the lamination of the double layer hydrophone was adopted in

bonding the multilayer structure of the imaging probe [6]. A pressing mechanism was

designed and developed, in which a “super ball” of about 4.6 cm in diameter was used to

press the layered structure of PVDF film during bonding. In this method, the hand

operated press was used to press the two layers of film in order to obtain a thin glue layer

in such a way that the applied pressure makes the drop of epoxy, kept at the middle of the

film, spread radially. This ensured that the epoxy is spread radially away from the center

of the film so that no air pockets will be formed. The thickness and uniformity of the

bond line was optimized after conducting series of several bonding trials. The average

thickness of bonds in the range of 1-2 µm was achieved. The Figure 11 shows the

pressing mechanism while in use.

Figure 11: Pressing mechanism while in use

99

By using a small drop of non-conductive epoxy (RBC#3200, RBC Industries, Inc.), the

PVDF layer was glued with the backing materials and kept pressed by the pressing

mechanism overnight so as to cure the adhesive at room temperature. Care was taken to

achieve a thin and uniform layer of bonding. If attempt is made to make the glue too

thin, there is a possibility of epoxy being washed out. The surface to be glued was fully

cleaned and the epoxy was mixed very carefully to avoid producing bubbles.

The electrical leads were secured to the long grooves made along the backing and housed

in a stainless steel housing, which was also grounded, thus providing effective electrical

shielding before potting it with nonconductive potting compound. The other end of the

leads was soldered to a BNC connector through diode pairs. By using the conductive

silver epoxy (#2902, Tra-Con, Inc.), the front surface was connected to the stainless steel

housing, which was grounded. One single layer and two three layer transducers were

constructed. Figure 12 shows the parts in different stages of fabrication process and

Figure 13 shows the completed transducers.

100

Figure 12: Different stages of the fabrication process of single element transducer

Figure 13: Completed single element transducers

101

6.6 Characterization of Non-Resonant Single Element-Single Layer Transducer and Multilayer Barker Code Polymer Imaging Transducer

The criteria examined in characterizing the transducers included bandwidth, frequency

response and sensitivity. The parameter studied includes the epoxy bond thickness.

6.6.1 Electrical Characterization

The typical input admittance and phase spectra for the single layer and three layer

transducer measured in air are shown in Figure 14 to16.

Figure 14: Plot showing the variation of admittance magnitude and phase of single layer transducer.

Figure 15: Plot showing the variation of impedance magnitude and phase of single layer transducer.

Figure 16: Plot showing the admittance magnitude of 56 µm piezo film, one layer and three layer transducer.

0.000

0.005

0.010

0.015

0.020


Adm

ittan

ce M

agni

tude

in M

hos

70

72

74

76

78

Adm

ittan

ce P

hase

in D

egre

es

Admittance magnitude plot

Admittance phase plot

0

400

800

1200

1600


Impe

danc

e M

agni

tude

in O

hms

-80

-78

-76

-74

-72

-70

Impe

danc

e Ph

ase

in D

egre

es

Impedance phase plot

Impedance magnitude plot

0.000

0.010

0.020

0.030

0.040


Adm

ittan

ce M

agni

tude

in M

hos

Resonance frequency at 19.8 MHz

56 micron PVDF film

Single layer transducer

Three layer transducer

102

6.6.2 Experimental system

The experimental configuration is shown in Figure 17.

Tektronix 2430Digital Oscilloscope

CH 1 CH 2

HP 8116A, 50 MHzFunction Generator

ENIRF Power Amplifier 20 dB Gain

Stepper Motor

Preamplifier

T/R Control Circuit

Sync

50 ΩOutput

DigiplanMotor Controller

Probe under test

Stainless Steel plate targetat 1 cm distance

Deionized water bathat room temperature

Coaxial cable interconnect

Figure 17: Experimental setup for pulse echo measurement

103

The transducer was mounted on a manipulator and the measurement system was capable

of adjusting the transducer in three orthogonal axes. The transducer was used as a

transmitter-cum-receiver so that it functioned in the pulse-echo mode. A voltage-source

drive was obtained by using the 50 Ω output port of a 50 MHz HP (Agilent) function

generator and was used to generate a monocycle sine burst excitation over the frequency

spectrum of interest. The peak amplitude produced by the function generator without

connecting the transducer was measured with the oscilloscope set at 50 Ω coupling. The

transducer was pulsed into deionized water. A high impedance polished stainless steel

reflector kept at a distance of 1 cm was used as the pulse reflector. The single cycle

strain echo returned from the reflector induces electric field distribution in the transducer

via piezoelectric action. The output of the transducer was taken through the specifically

built T/R (Transmit-Receive) switch and amplified in the custom built preamplifier and

viewed by the Tektronix 2430 digital oscilloscope.

The pressure level depends on the position of the observation point in the acoustic field.

At the short axial distance from the transducer face, the transmitted and the echoed

signals were assumed to be proportional to the average pressure falling upon the surface

of the transducer [35]. In the case of both the technical and practical measurement point

of view, one cm separation distance was chosen to evaluate the performance of the

transducer, although much closer distance is preferred. Throughout the experiment, the

temperature was monitored and found constant with the variation of less than ±10C.

104

6.6.3 Pulse-echo response

The performance of the polymer transducers was evaluated by measuring the pulse-echo

response and the pulse spectrum. The pulse echo response was measured by recording

the signal received from a flat stainless steel reflector placed at a distance of 1 cm from

the transducer. The pulse-echo response was used to quantify the receive sensitivity of

the transducer. The pulse echo response was evaluated using two parameters. One was

the response amplitude, which was defined as the maximum value of the output voltage,

and the other is the response time tp, defined as the time duration between the times at

which the envelope of the peak amplitude is –40 dB. Figures 18 and 19 show the pulse

echo voltage received by the single layer and three layer Barker code transducers,

respectively developed in this work with the stainless steel reflector kept at 1 cm distance

from the transducer. From the time domain pulse echo responses, it may be seen that

higher sensitivity (about 2 fold increase) was achieved in the case of three layer Barker

code transducer compared to single layer transducer.

Figure 18: Pulse-echo response (in time and frequency domain) of single element, single layer transducer at 1cm depth in water for a monocycle sine burst excitation.

Horizontal scale: 100ns/div Vertical Scale: 20mV/div

Peak at 7 MHz

Pulse echo response of single layer transducer at 1 cm

-8

-6

-4

-2

0


Puls

e ec

ho a

mpl

itude

in d

B r

e pe

ak

105

Figure 20 and 21 show the pulse echo response of single and three layer transducers at

different depths in water, respectively. It was seen that the bandwidth of the transducer

decreased as the ultrasound pulse penetrates deeper.

Figure 21: Pulse-echo frequency response of single element, three layer transducer at different depths in water for a monocycle sine burst excitation

Figure 20: Pulse-echo frequency response of single element, single layer transducer at different depths in water for a monocycle sine burst excitation

Figure 19: Experimental pulse-echo response (in time domain) of a single element, three layer Barker code transducer at 1 cm depth in water for a monocycle sine burst excitation.

Horizontal Scale: 50ns/div Vertical Scale : 30mV/div Peak at 10 MHz

-20

-15

-10

-5

0


Puls

e ec

ho a

mpl

itude

in d

B r

e pe

ak

1 cm depth

8 cm depth

20 cm depth

-20

-15

-10

-5

0


Puls

e ec

ho a

mpl

itude

in d

B r

e pe

ak

1 cm depth

8 cm depth

20 cm depth

106

Figure 22 shows the pulse echo spectral response of a single layer and three layer

transducer at 1 cm depth.

From the spectral magnitude response, it may be observed that the polymer transducers

showed no discernible electrically excitable mechanical resonance in their spectra. The

results show that excellent wideband performance was achieved with the –6 dB

bandwidth extending about 2 decades of frequency [46]. It indicated a peak frequency of

7 MHz, center frequency of about 15 MHz and –6 dB fractional bandwidth between 2 to

28 MHz (176%) for the single layer transducer and a peak frequency of 10 MHz, center

frequency of about 12 MHz and –6 dB fractional bandwidth between 2.5 to 21 MHz

(154%) for the three layers Barker code transducer, respectively. Such bandwidth is not

Figure 22: Pulse-echo frequency response of single and three layer PVDF transducers at 1 cm depth in water for a monocycle sine burst excitation.

-14

-12

-10

-8

-6

-4

-2

0

0 10 20 30

Frequency in MHz

Rec

eive

d pu

lse

echo

sign

al in

dB

re

peak


Three layer transducer

1L 3LPeak frequency: 7 MHz 10 MHzCenter frequency: 15 MHz 12 MHz-6 dB bandwidth: 176% 154%

107

available using conventional, resonant transducer design. The extremely broadband

response is typical of non-resonant PVDF transducer, which has low impedance backing.

From the frequency response plot, a general trend may be seen that the center frequency

declines with increasing number of layers. The measured center frequency of the single

layer transducer was 15 MHz while the three-layer transducer was 10 MHz. The

difference in acoustic impedances of the glue layer and PVDF, and the thickness of the

glue layer are some of the causes for downshifting of the primary thickness mode

resonance of the transducer. The increased thickness of the glue layer is one of the

causes for the decreased sensitivity of the transducer

The effect of epoxy bond between the layers of PVDF film was examined. It was found

that thicker bond degraded the transducer performance. The thickness of epoxy bond

between the layers and between the backing and the layer stack must be thin to optimize

the performance. Figure 23 shows the frequency response of one single layer and two

three layer transducers. It may be seen that the frequency response of three layer

transducer number 2 is improved due to lesser glue thickness compared to three layer

transducer number 1.

If the thickness of the intermediate glue layers is maintained as very thin (~1/10 of the

shortest wavelength) and constant, then the transducer’s center frequency will remain

unchanged and its sensitivity will improve linearly with the number of active layers.

108

6.6.4 Diffraction correction

The acoustic field from uniformly vibrating transducer consists of two components

namely a plane wave and an edge wave [57]. The plane wave propagates with its

wavefront parallel to the transducer surface, and the edge wave radiates outward from the

transducer edges. The edge wave interferes with the plane wave. The calculation of

finding the average pressure on the receiver and relating it to the pressure of an ideal

plane wave on the surface of the radiator is called diffraction correction.

A.S. Khimunin has provided a table [58], which will relieve the necessity of carrying out

the calculation of the corrections by means of the known formulae. The table presents

Figure 23: Pulse-echo frequency response of one single layer and two three layer single element polymer transducers at 1 cm depth in water for a monocycle sine burst excitation.

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Frequency in MHz

Rec

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Three layer transducer I

Three layer transducer II

109

numerical values of the modulus (pp−

) of ratio of the average pressure to the pressure of

an ideal plane wave on the surface of a circular transducer propagating into an uniform

isotropic medium without absorption verses two independent variables namely, the

dimensionless distance S = 2

2ka

zπ and the wave parameter ka, where k is the wave number

(ω/c), a is the transducer radius, and z is the distance between the radiating and receiving

surfaces. The values provided in the table are for amplitude diffraction correction.

The procedure for incorporating the diffraction correction to the experimental data

consists of the following sequence: The wave parameter ka is determined for the given

radius of the transducer and appropriate frequency/wavelength. The column with the

nearest wave parameter value is found from the tables of diffraction corrections. Such

procedure would be faultless only in cases where the effective radius of the transducer is

equal to the geometric radius. Practically, the piezoelectric transducer will have effective

radius different from the geometric radius because of the disparity in the mechanical

tension between the periphery and the center of the source. So the correction effected

had some systematic error.

The diffraction correction is important at the lower frequencies and it decreases with

increasing frequency. Hajime Seki et.al. has pointed out that one decibel per distance of

λ

2a provides a rough criteria and a general estimate of the attenuation due to diffraction

and they have given the expression for the “attenuation due to diffraction” (αd) as [59]:

110

αd = fa

c27.1 dB/cm. (6.3)

The Figure 24 shows the pulse echo response (both in time domain and frequency

domain) of a single element, single layer transducer at different depths for a monocycle

sine burst excitation. It is seen that the spectrum of pulse varies as it penetrates deeper

into water because the attenuation in water/tissue is frequency dependant. It also

revealed that the center frequency and bandwidth of the ultrasonic pulse decrease as the

ultrasound pulse penetrates deeper. This implies that the axial resolution of the beam will

worsen as the ultrasound beam penetrates deeper into tissue unless time gain

compensation (TGC) is applied in clinical use. Moreover, the results show that at close

distance, more or less unipolar pulses are produced. With increasing separation the

pulses degrade towards bipolar behavior. The unipolar degradation observed at close

distance is due to the loss of low frequency content because of diffraction losses. The

consequence of this diffraction loss is a degraded signal/noise ratio at the lower end of

the frequency spectrum.

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At 6 MHz

At 20cm depth

At 1 cm depth

Figure 24: Pulse echo response (in time and frequency domain) of single element, single layer polymer transducer at different depths in water for monocycle sine burst excitation

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The diffraction loss corrections were applied to the measured frequency spectra in the

pulse-echo mode. The Figure 26 illustrates the frequency spectra of a single layer

transducer with and without diffraction correction. It may be seen that there is an

improvement in the low frequency arena and consequently there is a change in the

estimated –6 dB bandwidth due to the diffraction effects.

Figure 26: Pulse echo response of a single layer transducer at different depths in water for a monocycle sine burst excitation (with and without diffraction correction).

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0.6

0.7

0.8

0.9

1

0 5 10 15 20 25 30 35 40Water path length in cm

Nor

mal

ized

ech

o si

gnal

Figure 25: Pulse echo response of a single element, single layer transducer for different water path length at peak frequency for monocycle sine burst excitation.

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1 cm depth - without diffraction correction

1 cm depth - with diffraction correction

8 cm depth - without diffraction correction

8 cm depth - with diffraction correction

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6.6.5 Sensitivity correction

In the receive mode, the active element of the PVDF transducer acts like a voltage source

in series with a small capacitor which depending upon the construction of the transducer,

will vary. The source is loaded by the cable from the active element to the measurement

device (Oscilloscope). So the receive voltage that obtained is in term of effective

voltage. In order to know the exact voltage produced by the active element of the

transducer at its terminal, correction for the capacitance loading can be performed.

The schematic diagram of the transducer acting as a voltage source is shown in Figure 27.

Probe Coaxial interconnect cable Preamplifier Oscilloscope

Ce

VocCs Cc CL

Voc Vec Veff

Figure 27: Schematic diagram showing the capacitance loading of the polymer transducer

114

Voc represents the open circuit voltage

Vec represents the end of cable voltage

Veff represents the effective voltage measured.

The voltage source Voc, produced by the active element of the probe having a capacitance

Ce is driving the coaxial cable with a capacitance of Cc, the stray capacitance, Cs and the

measuring apparatus (Oscilloscope) with a capacitance of CL.

The open circuit voltage of the active element of the transducer is estimated as [31]:

Vmeasured = Voc

++ SICT

T

CCCC (6.4)

where

CT is the capacitance of the transducer element

CIC is the capacitance of the interconnecting cable

CS is the stray capacitance

The change in sensitivity caused by the change in the capacitive loading between the

transducer element and the measurement device is a factor of

++ SICT

T

CCCC . It may

be seen that a change in the interconnecting cable capacitance would directly affect the

sensitivity of the transducer.

6.7 Summary

The single element, single layer transducer and multilayer Barker code transducers were

developed. The experimental results indicated that the polymer transducer could be

operated in the clinically relevant frequencies using a single transducer. The pulse echo

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spectral response of a single layer and three layer transducer indicated a –6 dB fractional

bandwidth between 2 to 28 MHz (176%) for the single layer transducer and 2.5 to 21

MHz (154%) for the three layer Barker code transducer, respectively.

It has been reported in literature that the temperature dependence of the PVDF transducer

performance for the –40 to 800C range showed an approximate linear decrease in center

frequency and increase in fractional bandwidth with increasing temperature. Since the

speed of sound decreases approximately with increasing temperature, PVDF transducer

would have a temperature dependent peak in its performance. The water temperature was

monitored during the period of testing and the variation of temperature was ±10C during

the measurement. The measurements were carried out at room temperature.

6.8 Conclusion

It is expected that this wideband non-resonant transducer will become a useful clinical

tool with expanded diagnostic ability and minimize the trade off between bandwidth and

penetration depth. This wideband polymer transducer, once optimized, will be useable in

all clinically relevant frequencies between 2 to 15 MHz using a single transducer and will

improve the diagnostic efficacy. This single element transducer will be well suited for

high frequency imaging such as ultrasound backscatter microscopy. High frequency

single element transducers are currently used only in some specific applications such as

in the field of ophthalmology and dermatology. Since most of the diagnostic applications

use multielement transducers, it was decided to explore the feasibility of implementing

the Barker code concept in a linear array transducer.

116

The next chapter deals with the design and development of a multielement (array), single

layer transducer and a multilayer transducer in which the Barker code concept used in

single element was extended to multielement structure as a feasibility study.

117

CHAPTER 7: DESIGN, DEVELOPMENT AND CHARACTERIZATION OF NON-RESONANT MULTIELEMENT (ARRAY), SINGLE

LAYER AND MULTILAYER POLYMER IMAGING TRANSDUCERS

This chapter deals with the design and development process of multielement transducer

incorporating the concept of Barker code and the results obtained as a part of feasibility

study.

7.1 Introduction

The properties and techniques developed in the preceding sections on single element

transducer were applied to multielement (array) transducer. Several engineering

approaches such as electrode patterning design, gluing and lamination techniques,

assembly procedures and mounting arrangements were adopted in the array development.

7.2 Design Consideration and Description of Array Structure

The multilayer, multielement linear array transducers were designed and developed by

stacking 56 µm thick PVDF film according to Barker code. Design requirements

imposed by the intended clinical application and system electronics normally determine

the number, dimensions, and spacing of the array element. The length (L) of single

element of the array is generally restricted by anatomical considerations; typical values

range from a few mm to few cm. The frequency response of a single element is the

Fourier transform of the excitation voltage across the aperture. As the element width

increases, the Fourier Transform narrows and the element response decreases more

rapidly with angle [10]. The design of a transducer depends on the depth of view

118

(penetration) imposed by the application and the targeted organ and the footprint

(aperture) of the transducer. The aperture describes its lateral dimensional area. For

higher frequencies, smaller aperture is preferred in order to focus at shallow depth.

Theoretically, the element width (W) of an array should be less than or equal to half

wavelength in water to avoid grating lobes in imaging. The array should have large

number of elements in order to obtain a high-resolution image. Arrays having elements

ranging from 32 to 512 are currently available [9]. If the number of elements is reduced,

the quality of constructed images may be lowered. There are trade-offs between the

quality of constructed images and the number of elements (or the element spacing) of an

array. The probe output response should have Gaussian like shape for good imaging.

Four element array design was chosen in this work since a group of four elements are

usually fired together in the case of imaging arrays having larger elements in practice.

Figure 28 shows the schematic view of the array element and bonding pad design at one

side of a PVDF film of the linear array transducer.

119

It is a 4-element device with element widths (W) of 1.7 mm and an inter element spacing

of 0.3 mm and total length (L) of 10 mm. The array aperture is 7.7 mm in azimuth and

10 mm in elevation. This approximates a square transducer source. The electrode pattern

on opposing surface of the film is so defined in such a way that by laminating the films,

multilayer structure is created. It has a unique design of bonding pads specifically

suitable for implementation of Barker code pattern for a three-layer four element linear

array. The bonding pads are spread on all four sides of the film. The bonding pads have

dimensions of 1.5*1.5 mm, which facilitated making direct epoxy (Tra-Con 2902)

connection of the pads to the electrical leads. Every element in the array has one

interconnect for the electrical signal and another for ground. To reduce the number of

vias in the multilayer structure, via was shared by the adjacent layers by capacitive

Figure 28: Schematic representation of array element and bonding pad pattern at one side of the polymer film of the proposed linear array transducer.

(Not to scale)

1. 5

1. 5

120

coupling. The unelectroded area of the film prevented the shorting of the signal electrode

to ground wire and vice versa. The electrode legs of each element on each side of the

three films of the proposed three layer transducer are designed in such a way that they lie

at different locations of the four sides and not overlapping in any area to avoid possible

electro-acoustic and capacitive coupling between each element of the array and to the

adjacent layer.

7.3 Fabrication Process

Most of the processes were developed by using several logical sets of investigative steps.

For example research was done to determine the candidate method and a test method was

adopted and the best method was implemented. The electrode pattern (on one of the

sides of the layer) for developing the three layer Barker code array transducer shown in

Figure 28 was used in fabrication of the four element array. The Figure 29 shows the

photograph of designed electrode pattern (on one side of one of the layers).

121

The sputtering mask was made by “Electro Discharge Machine” (EDM) according to the

designed array pattern by using a slightly curved thin stainless sheet of 0.012” thick,

which facilitated in holding the poled polymer film tightly when the mask was pressed

against the film that resulted in good edge definition of the electrode during sputtering

process.

The electrode and electrical leads were made on the uniaxially stretched polarized PVDF

film of 56 µm thick by vacuum deposition by using the metallic mask by the local

company, Measurement Specialty Inc., Norristown, PA. The electrodes were sputtered

onto the faces of the polarized PVDF films with 200 Angstrom thick of Ni (V)

Figure 29: Actual electrode pattern of an array layer

122

vanadium (7%) mixed Nickel (Vanadium was used to make the alloy non magnetic) as

base metal and 800 Angstrom thick of gold (Au) on top of the Nickel which serve as

electrode. Before sputtering, the piezofilms were given thermal and corona treatment for

thermal stability and good metal adhesion, respectively.

The leads were connected permanently by means of a conductive silver epoxy (Tra-Con

2902). A novel method of array stack architecture employing the Barker code design was

used. The stacking and bonding technique used in developing the single element

multilayer transducer was adopted in fabricating the multielement linear array transducer.

A non-piezoelectric polymer rod of square shape made from (Kynar) polyvinylidene

fluoride (PVDF) resin and was machined and polished smoothly on the gluing side and

roughened on the other end to prevent acoustic reverberation. Polymer backing was

chosen in order to obtain fundamental thickness resonance of half wavelength. Mating

grooves were made on the sides of the backing to secure the leads. Several trial pressings

were conducted in order to optimize the uniformity and thickness of the glue layer. The

PVDF films were bonded together by using a slow curing, low viscosity epoxy (RBC

3200) and by using the pressing mechanism shown in Figure 11. Pressure was applied

during bonding in order to achieve a glue line of less than 2 µm thick. The bonded

assembly was kept in the pressing mechanism overnight until the epoxy was cured at

room temperature. After bonding the films on to the backing, the entire structure was

placed in the stainless steel housing. Figure 30 shows the main components of the array

at different stages while construction and Figure 31 shows the completed array

transducers.

123

Figure 30: Different stages of the construction process of four element array transducer

Figure 31: Completed array transducers.

124

7.4 Design and Development of System Electronics and Control

A complete transmit system for the transmit circuitry and an operational amplifier for the

receive circuitry were built in order to in evaluate the array transducer in pulse echo

mode. Figure 32 shows the transmit/receive switching box and the preamplifier.

7.5 Experimental System for Performance Evaluation

The experimental arrangement used was similar to that described earlier and used for

testing the single element wideband PVDF transducer. Figure 33 shows the water tank

experimental system setup used to measure the pulse echo response of the array

transducer.

Figure 32: Photograph of the newly built transmit/receive control circuit and the preamplifier

125

By employing the specially designed and fabricated dual transmit/receive (T/R) switch,

each element in the array could be connected both to the pulsing unit through diode pairs

and to the receiving unit (Oscilloscope) through the custom designed preamplifier. In

this way an individual element of the array and all four elements of the array could be

energized and tested in the pulse echo mode. The element was excited with a monocycle

sine burst. The echo signal was reflected from a stainless steel reflector kept at 1 cm

depth (to reduce attenuation) away from the array in the water, fed into the corresponding

element of the array and was amplified by the preamplifier before sending it to the

Tektronix 2430 digital oscilloscope.

Figure 33: Experimental configuration for pulse echo measurements

126

7.6 Performance Evaluation of the Array Transducer

The ultimate test of an array is the ability to obtain a good image. In an ultrasound

imaging system, the performance of the transducer is generally masked by system

electronics used to drive the array elements. Therefore the evaluation of individual

elements is the best benchmark to assess the performance of any array. The criteria

examined in evaluating the performance of the transducers fabricated in the course of this

research include bandwidth, frequency response, and uniformity and pulse echo

sensitivity. This section describes the experimental measurements performed on the

array transducer.

7.6.1 Single layer array transducer

7.6.1.1 Pulse-echo response of individual element

The pulse echo experiment was designed in order to test the performance of the array.

The time domain response of each element in pulse echo mode is the widely used

measure in array performance. The pulse echo response was measured in order to

determine the bandwidth, sensitivity, pulse length and center frequency of each element

of the array. Each element of the array transducer was excited with a monocycle sine

burst and the pressure wave, reflected from the stainless steel reflector placed in the water

tank 1 cm away from the face of the array, was received by the same element and fed into

the receiving circuit. A custom built preamplifier with a gain of 6 dB was used in order

to have a satisfactory display of the echo amplitude on the oscilloscope. The highest

peak-to-peak amplitude of the echo signal was recorded over the frequency range of 1 to

40 MHz, and the center frequency, -6 dB bandwidth, sensitivity and the pulse length were

127

determined. The highest amplitude was used to assess the sensitivity of the measured

element. The pulse length was determined from the first and last points of the waveform

that were 40 dB down in amplitude relative to the peak [60].

Figure 34 shows the experimental frequency response recorded for one of the elements of

the single layer array, which is representative of a typical response observed during

testing. The response indicates a –6 dB bandwidth of 155% (between 4 to 32 MHz).

At 14 MHz Horizontal Scale: 50ns/div Vertical Scale: 10mV/div

Figure 34: Experimental pulse-echo responses (in time & frequency domain) for one of the elements of single layer array transducer at 1cm depth in water for a monocycle sine wave excitation

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128

7.6.1.2 Pulse echo response of all the four elements

All four element of the array were excited simultaneously and the pulse echo response

was measured. This is a typical case in reality in medical imaging, as a group of four

elements are usually fired at a time and then sequentially advanced further in the case of

arrays having larger number of elements. Figure 35 shows the experimental pulse echo

response.

Horizontal Scale: 50ns/div Vertical Scale : 20mV/div

Peak at 12 MHz

Peak frequency: 12 MHz Center frequency: 17.5 MHz -6dB bandwidth : 166%

Figure 35: Experimental pulse-echo responses (in time & frequency domain) while exciting all the four elements of single layer array transducer at 1cm depth in water for a monocycle sine wave excitation

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129

The pulse echo response of all four elements of single layer array transducer when

excited together indicated a –6 dB fractional bandwidth of 166% (between 3 MHz to 32

MHz). From the time domain signal, it may be seen that the transducer radiates a clean

acoustic wave almost without any ringing. This is desirable as shorter impulse response

implies shorter acoustic pulse, which allows better axial resolution in ultrasound imaging.

7.6.1.3 Uniformity

The pulse echo response of the individual elements was also examined. A uniform

ultrasonic response of individual elements throughout the array was observed. A typical

recorded time domain response of one of the elements of the single layer transducer is

shown in Figure 36 (Fig.34 reproduced for convenience). The arrays constructed using

PVDF were having uniform acoustic response from element to element. The waveforms

for all the elements were virtually identical and the response amplitude typically varied

within ±1 dB. This property is important for the array to produce high quality ultrasound

images.

At 14 MHz Horizontal Scale: 50ns/div Vertical Scale: 10mV/div

Figure 36: A typical observed time domain pulse echo response of one of the elements of the single layer array transducer at 1cm depth in water for a monocycle sine burst excitation.

130

7.6.2 Multilayer array transducer

7.6.2.1 Pulse echo response

Each stack of elements of the three layer array transducer was excited with a mono cycle

sine burst and the pulse echo response from the stainless steel reflector placed at 1cm

depth in water was measured. Figure 37 shows the experimental pulse echo response of

one of the stacks of the three layer transducer. It shows a –6 dB bandwidth of 141%

(between 4 to 23 MHz)

Figure 37: Pulse-echo responses (in time & frequency domain) of a representative element stack of three layer array transducer at 1cm depth in water for a monocycle sine wave excitation.

Horizontal Scale: 50ns/div Vertical Scale: 20mV/div

Peak at 10 MHz

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Figure 38 depicts the time domain pulse echo response of single and three layer array at

similar excitation condition. It may be seen that higher sensitivity (about a 2 fold

increase) was achieved in the case of three layer transducer compared to a single layer

array.

Peak at 14 MHz Horizontal Scale: 50ns/div Vertical Scale: 10mV/div

Horizontal Scale: 50ns/div Vertical Scale: 20mV/div

Peak at 10 MHz

(A)

(B)

Figure 38: Typical observed time domain pulse echo response of one of the elements of single layer (A) and one of the element stacks of three layer (B) array transducer at the peak frequency at 1cm depth in water for a monocycle sine burst excitation.

132

Figure 39 shows the over layed pulse echo response of two stacks of elements of the three

layer array transducer, which demonstrated the excellent uniformity in the pulse echo

response.

7.7 Summary

Recording the spectral response of echoes from the element stack of the array it was

found that the experimentally achievable fractional bandwidth was 155% in the case of

single layer polymer array and 141% in the case of multilayer Barker code array

transducer. All the elements of the single layer array displayed greater than 150%

bandwidth, and the element stack of three layer transducer displayed a bandwidth greater

than 140%, which is one of the design goals of this research. The three layer array

exhibited increased pulse echo sensitivity compared to single layer transducer.

Figure 39: Pulse echo frequency response of two element stacks of three layer array transducer at 1cm depth in water for a monocycle sine burst excitation.

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7.8 Conclusion

Single layer and multilayer array transducers have been constructed. The construction of

multilayer array required critical alignment during bonding. The electrical connection

was one of the challenging tasks. The array has been tested in the pulse echo mode and

the bandwidth, sensitivity and uniformity were examined. The bandwidth and uniformity

of the elements/element stack were found to be excellent and the sensitivity of the three

layer array was found to be higher than that of the single layer array. This simple array

design gave a satisfactory performance and extension to arrays having large number of

elements would be straightforward. It is expected that this kind of array probe will find

application in practical imaging systems. A full understanding of the performance of the

newly designed and developed array would require extensive modeling. However, the

results presented here will help for future work

The next chapter presents a consolidated summary of the research work accomplished,

conclusions and suggestions for future work based on the outcome of this work.

134

CHAPTER 8: SUMMARY AND CONCLUSION

8.1 Summary of the Research Work

The simultaneous improvement of both the sensitivity and bandwidth, in the case of

imaging transducers and hydrophone probes has been explored in this thesis. The

research work presented provides useful information on the performance of PVDF probes

for medical ultrasound application. The results presented in each of the preceding

chapters lay important ground work for further advancement in employing PVDF

technology and incorporating the Barker code concept to the imaging transducer design.

The basic ultrasound imaging requirements, material properties and selection, design

criteria, fabrication process, and the performance evaluation procedure were described in

detail.

Double layer hydrophone probes using dissimilar thickness of PVDF film were fabricated

and the probes parameter such as sensitivity, frequency response, bandwidth, angular

response, effective aperture size and orientation (in two planes) effects were determined.

This design exhibited simultaneous enhancement of sensitivity and bandwidth. The

development of this type of polymer hydrophone probe using dissimilar thickness of

PVDF film was the first one of this kind made and commercially available and no such

development was reported in the literature as to the best of the knowledge of this author.

135

To address the issue of non-existence of calibration standard below 500 kHz and the need

for a new procedure, a measurement method employing planar scanning technique was

developed to determine the frequency response of double layer polymer hydrophone

probe below 1 MHz. The method developed could also be used while characterizing the

newly developed ultrasonic transducers for medical imaging for accurate measurement of

peak rarefactional pressure used in predicting the potential biological effects in

ultrasound imaging. This method using planar scanning technique for the low frequency

calibration of hydrophone has not been explored by any one previously as to the best of

the knowledge of the author.

Transmit and receive electronic systems for the single element and multielement

transducer were designed and developed.

Single element, single layer transducer and multilayer imaging transducers using Barker

code concept have been fabricated and the performance of the multilayer was compared

with the single layer in terms of bandwidth and sensitivity. The results show that

excellent wideband performance was achieved with the –6 dB bandwidth extending about

2 decades (~20 MHz) of frequency. It indicated a –6 dB fractional bandwidth of 176%

(between 2 to 28 MHz) for the single layer transducer and –6 dB fractional bandwidth of

154% (between 2.2 to 21 MHz) for the three layer Barker code transducer respectively.

Such bandwidth is not available using conventional, PZT transducer design. If a

commercial version of the enhanced bandwidth imaging transducers were available, an

effective trade-off between the penetration depth and resolution capability could be

136

achieved in clinical practice as the operator will be able to control the frequency with out

changing the imaging transducer. The enhanced bandwidth would also allow the

transducer to be used for fundamental and higher harmonics imaging.

As single element transducers are seldom used at clinically relevant frequencies between

2 – 15 MHz, a simple four-element array was designed with 56 µm thick PVDF film to

test the feasibility of a more desirable transducer implementation. Single layer array

transducer and multilayer Barker code array transducers were developed and the

performance was evaluated in pulse echo mode. The results indicated a fractional

bandwidth of about 155% (between 4 to 32 MHz) in the case of single layer polymer

array. All the elements of the array displayed greater than 150% bandwidth and showed

very good uniformity of impulse response. The three layer Barker code array transducer

was also tested in pulse echo mode and exhibited a fractional bandwidth of 140%

(between 4 to 23 MHz) and excellent uniformity. The time domain pulse echo response

of three layer array transducer revealed increased sensitivity (about 2 times) compared to

a single layer array transducer.

8.2 Conclusion

The work presented in this thesis has concentrated maximizing the bandwidth of

piezoelectric transducers for medical ultrasound applications utilizing PVDF. The work

carried out has shown that the single element and multielement (array) transducers

constructed using PVDF demonstrated excellent operating characteristics such as broad

bandwidth and uniformity of response. The relatively low pulse echo sensitivity of

137

PVDF polymer transducer was partly overcome by adopting the Barker code technique.

The pulse echo response of the transducers was studied, which is an important aspect

directly relevant to clinical application of ultrasound imaging. The imaging transducer

developed using piezopolymer has demonstrated excellent bandwidth with simultaneous

improvement in pulse echo sensitivity. One of the problems in fabricating the multilayer

transducer was the uniformity and thin adhesive layer in bonding the polymer film. Still

more process control is needed on the quality of epoxy adhesion and uniformity of glue

layer in order to improve the performance of the transducer.

The transducer design described could be operated at frequencies well below

fundamental thickness mode resonance of a single layer and could operate in the whole

frequency range 2-15 MHz, which is relevant for current clinical applications.

Consequently, one transducer could be sufficient for almost all diagnostic imaging

applications. Such a wide bandwidth is not available using conventional resonant

transducer design. Once fully developed, such an enhanced bandwidth transducers would

minimize the trade off between bandwidth and penetration depth and shorten examination

time, allow immediate, on-site optimization of image resolution and therefore lead to

improvement in diagnostic efficacy. The literature search revealed that no one has

developed the Barker code multilayer PVDF array transducer so far.

It has also been shown that the use of PVDF in imaging transducers provides enhanced

bandwidth without the need for quarter wave matching layer as required in the case of

PZT material. The good acoustic impedance matching of piezopolymer with human

138

tissue and water makes the polymer a good candidate in efficient energy transfer from the

imaging transducer into the propagation medium. Since polymer materials inherently

possess low mechanical quality factor (Qm), broader bandwidth and short impulse

response can be achieved, which results in better image resolution and flexibility in

optimizing the image of structures at various depths. Since the piezopolymer is a flexible

material, it is easy to fabricate probes with curved surfaces. However, due to low

electromechanical coupling coefficient and low dielectric constant, the piezopolymer is

disadvantageous compared to piezoceramic in terms of sensitivity and transmission

efficiency. By employing the multilayer structure using Barker code concept, it has been

demonstrated that the sensitivity could be improved. The polymer transducer has the

inherent drawback in electrical impedance match with the pulser due to its low dielectric

constant. However, this can be resolved to some extent by electrical tuning. In such

cases, we have to sacrifice the bandwidth to some extent. Anyhow, as long as we obtain

a good sensitivity and reasonable bandwidth useable in clinical range, piezopolymer will

be a good candidate for ultrasound imaging applications.

8.3 Suggestion for Future Work

The development of PVDF polymer ultrasonic transducers is an exciting area of research.

The experience gained in the course of development of multilayer polymer transducer

suggests that the design presented could be further improved. The alignment of layers,

use of some other lower viscosity bonding adhesive, the thickness of the bonding

adhesive, and by designing a viable way to implement electrical connections to each of

the individual array elements in the multilayer transducer design. This is a challenging

139

task as a typical imaging array has more than two hundred elements. Placing the transmit

and receive electronics circuitry and the preamplifier closest to the array element needs to

be explored in order to improve the sensitivity still further so that the array element will

drive the preamplifier instead of driving the connecting coaxial cable. An integrated

design approach in developing the transducer and its compatible pulsing and receiving

units taking into consideration of impedance matching is worth considering. Finally in

the author’s opinion, innovation by combining the single crystal and piezopolymer

material (may be designated as polymer /crystal composite) will be foreseen.

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[61] “Characteristics and calibration of hydrophones for operation in the frequency range 0.5MHz to 15MHz”, International Electrotechnical Commission, Geneva, Switzerland, Publication IEC 60866, 1987. [62] B. Zeqiri and A.D. Bond, “The influence of waveform distortion on hydrophone spatial-averaging corrections – Theory and measurement”, J.Acoust. Soc. Am. vol. 92(4), pp.1809-1821, October1992. [63] G.R. Harris and D.G. Shombert, “A pulsed Near-Field Technique for Measuring the Directional Characteristics of Acoustic Receivers”, IEEE Trans. UFFC, vol. 32(6), pp. 802-808, 1985. [64] P. Lum, M. Greenstein, C. Grossman and T.L. Szabo, “High-Frequency Membrane Hydrophone”, IEEE Trans. UFFC, vol. 43(4), pp.536-544, July 1996. [65] “Acoustic Output Measurement and Labeling Standard for Diagnostic Ultrasound Equipment”, AIUM Standards Publication, 1992. [66] L.E. Kinsler, A.R. Frey, A.B. Coppens and J.V. Sanders, Eds. Fundamentals of Acoustics. New York: 3rd ed., John Wiley, 1982, p.107. [67] D.G. Shombert, S.W. Smith, and G.R. Harris, “Angular response of miniature ultrasonic hydrophones”, Medical Physics, vol. 9(4), pp. 484-492, July/Aug. 1982. [68] G.R Harris, “Hydrophone Measurements in Diagnostic Ultrasound Fields”, IEEE Trans. UFFC, vol. 35(2), pp. 87-101, March 1988. [69] D.R. Bacon, “Properties of a PVDF Hydrophone with 100MHz bandwidth for studying Medical, Non-linear and other Fields”, Ultrasonic Symposium 1980, pp. 582-585.

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APPENDIX A: DEVELOPMENT AND CHARACTERIZATION OF IMPROVED

DESIGN OF DOUBLE LAYER POLYMER HYDROPHONE

PROBE

a. Introduction

In diagnostic ultrasound imaging, there is always concern about the collateral tissue

damage due to exposure to ultrasound. The ultrasound device sometimes affects the

efficacy of patient treatment and possible side effects. The ultrasound field produced by

the newly developed imaging transducers need to be characterized by performing

acoustic measurements using a device called hydrophone as recommended by FDA [11],

AIUM/NEMA [12], and IEC [61] before use in clinical practice. It measures the spatial

distribution of the acoustic field (including its size, shape and energy) produced by the

medical imaging transducers and it should provide reliable measurements and comply

with the requirement of international standards. Hydrophones are made using either

piezoceramic or piezopolymer. The piezoceramic hydrophones have large physical

dimension and inadequate bandwidth resulting in temporal distortion and spatial

averaging of the waveform. Hence they are not suitable for complete quantitative

measurements of medical ultrasonic fields. In view of its small sizes of sensing element,

inherent broadband properties, acoustic transparency, and history of successful usage in

characterizing the ultrasonic field produced by medical imaging transducers, polymer

hydrophones are preferred and accepted for evaluating the imaging probes.

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The primary criteria for such a polymer hydrophone are that the hydrophone should

possess adequate sensitivity and broad bandwidth in order to determine the temporal

variation of the pulse waveform accurately during the characterization of diagnostic

ultrasound transducer. Since the sensitivity of the hydrophone depends on the voltage

produced by the active element, thinner element will have lower sensitivity, as the active

volume of the piezoelectric material is smaller. The sensitivity could be improved by

having a large size of sensing element, since this increases the hydrophone’s capacitance.

But if the hydrophone’s active element is large compared to the beam dimension, it will

introduce pressure-averaging effect over the active element causing under estimation of

the acoustic pressure and result in error in the derived acoustic quantities such as Ispta, MI

etc. Although numerical corrections to the spatial averaging effect could be made [62] to

compensate for the error caused, such corrected acoustic parameters from the actual

measured data will not be accurate. So the active element size could not be increased in

order to obtain improved sensitivity.

However, the sensitivity of the hydrophone could be increased by increasing the

thickness of the PVDF film but the bandwidth will be reduced by using thicker film, as

the resonance frequency is inversely proportional to the thickness of the PVDF film.

Harris [63] has discussed the necessity of having wider bandwidth hydrophones in order

to characterize the imaging transducers generating higher harmonics through non-linear

propagation in water. He pointed out that the acoustic parameters derived from the data

measured using hydrophones with inadequate bandwidth showed large error (exceeding

30%) in determining the value of Mechanical Index (MI) [24].

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Double layer PVDF hydrophones are conventionally made using two similar thickness of

PVDF film (25+25 µm and 9+9 µm) and have the measured fundamental thickness

resonance of approximately 20 and 50 MHz respectively. The thickness of the two layers

determines the resonance frequency of the hydrophone (fr = c/λ = c/2t, where fr is the

resonance frequency, c is the speed of sound in PVDF, t is the thickness of the two layers

of film and λ is the wavelength). Thinner film provides higher resonance but shows

comparatively lower sensitivity for a given spot size. The double layer polymer

hydrophone design using similar thickness of PVDF film was originated at the National

Bureau of Standards [15]. Hydrophone using dissimilar thickness of PVDF film (25+9

µm) was developed incorporating better electrical shielding and cable termination.

In order to have a robust device that will provide simultaneous improvement of both the

bandwidth and sensitivity, a new approach was adopted in this work, which involved

laminating a thicker active PVDF film with a thinner ground film, while keeping the

standard element size of 0.4 or 0.6 mm diameter. It is worth mentioning here that the

development of the double layer membrane hydrophone using different film thickness

(25+9 µm) at the beginning of the year 1999 was the first one of such kind, which

demonstrated enhanced bandwidth compared to that of hydrophone made with 25+25 µm

thick PVDF film. The polymer hydrophones developed using this improved design were

sold by my former employer (Perceptron, Inc.) to several researchers and manufacturers

of medical ultrasound equipment around the world for use as a reference hydrophone to

perform on the site calibration of other hydrophones and also as working hydrophones for

accurate acoustic measurements. My former employer (Perceptron, Inc) was the only

150

commercial manufacturer of double layer membrane hydrophone in USA and second in

the whole world (the other manufacturer being “Marconi” in UK). Fabrication of several

hydrophones with similar thickness of PVDF film has provided me the opportunity in

producing a hydrophone with enhanced performance by using dissimilar thickness of

PVDF film.

Several PVDF double layer hydrophones with similar thickness (25 µm+25 µm and

9 µm+9 µm) and dissimilar thickness (25 µm+9 µm) of PVDF film were developed. In

order to provide the physical basis of the behavior of the hydrophone, experimental

determination of the different properties were investigated. Besides determining the

frequency response, bandwidth and sensitivity, the directional response was also

measured and the variation of the effective spot size of the hydrophones having different

geometrical spot size, different film thickness, were studied at different frequencies and

at different orientation.

b. Background

In order to reproduce the acoustic pressure faithfully, the effective diameter of the

hydrophone should be smaller than the acoustic wavelength. In measurement practice,

when a membrane hydrophone is used in continuous wave (CW) or long tone-burst

excitation conditions, acoustic reflections will occur at the membrane surface allowing a

standing wave to develop. To eliminate the standing wave, the hydrophone is tilted so

that the source transducer and the hydrophone are not in parallel. By doing this, the

received hydrophone signal is decreased, as the voltage response of the hydrophone is

151

dependent on the angle at which the ultrasound wave is incident on the hydrophone.

Hence the angular response (the dependence of the amplitude sensitivity of hydrophones

upon the incident wave front direction) of the hydrophone is to be known to assess the

error introduced and to make appropriate directivity correction. Further, the knowledge

of the effective diameter of the hydrophone is important in assessing the spatial averaging

effect of the hydrophone [63]. This is a routine measurement practice while measuring

the acoustic output of ultrasonic transducers, particularly when using the higher

frequency transducers that are currently used in diagnostic imaging (2-15 MHz).

Some ultrasound imaging applications demand higher frequency transducers to image at

closer distances. Although conventional diagnostic imaging transducers operate from 2-

15 MHz range, very high frequency (VHF) transducers are used for eye, skin,

intraoperative and endoluminal applications. For instance, for a transducer having 1 cm

diameter and operating at 20 MHz, the –6 dB beamwidth at 1cm distance would be 0.08

mm. The geometrical diameter of the hydrophone commonly used in acoustic

measurement of ultrasonic probes is 0.4 mm. Since the diameter of the hydrophone is

large compared to the –6 dB beamwidth, spatial averaging will take place resulting in

substantial underestimate of the true acoustic pressure during measurements. To alleviate

this problem, peak pressure amplitude at the acoustical axis and at a distance of one

hydrophone radius off-axis is measured at all the four directions from the acoustical axis.

From the measured data, the true on-axis acoustic pressure is estimated by applying

corrections to the measured acoustic parameters [62]. In order to perform this

measurement, the effective diameter of the hydrophone needs to be known preciously.

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Moreover, the actual sensing element of the hydrophone will not be exactly circular,

although the physical construction is circular in shape. There will be asymmetry in the

shape due to fringe field effect during electrical poling of the hydrophone probe.

Therefore, the exact active diameter of the hydrophone probe and its corresponding

orientation setup as used while determining the effective diameter are to be known for

accurate measurement of acoustic pressure radiated by the ultrasound transducers.

To investigate the variation in effective diameter of hydrophones and to predict the

directional behavior and dependency on frequency, amplitude directivity measurements

were conducted by employing several PVDF membrane hydrophones of different kind,

various geometrical spot size and dissimilar thickness of PVDF film in two orthogonal

axes of the hydrophone. The angular response and the effective diameter of the

hydrophone are usually determined from the measurement of its directivity pattern. The

directivity pattern was obtained by rotating the hydrophone in the far field of the

transducer's plane wave, and by measuring the response at some angle of rotation θ and

plotting the response as a function of θ [31]. From the directivity patterns, the effective

diameter of the hydrophone is determined (details provided in the later section). In order

to ascertain the sensitivity and bandwidth, the newly fabricated hydrophones were

evaluated by using Time Delay Spectrometry (TDS) technique combined with

substitution method [2, 3, 4].

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c. Physical Description

The polymer hydrophone was made from acoustically transparent PVDF material with a

small central portion made piezoelectrically active. Hence, the hydrophone being an

unbacked membrane, operates in the thickness extensional mode and the fundamental

resonance frequency is determined by the thickness of the PVDF film. One of the two

layers has a poled central circular positive electrode on one side and the negative ground

plane on the other side. The second layer does not have any active element but simply

provide protection to the active element besides reinforcing the grounding. The standard

double layer hydrophones usually have active diameter of 0.4 or 0.6 mm, with and

without preamplifier. If preamplifier is not included, then the coaxial cable length is

usually about 75 cm. The hydrophone has an overall dimension of 11 cm in diameter

with the exposed membrane having 7 cm in diameter.

d. Design Consideration

It is essential that the hydrophone is capable of capturing both the compressional and

rarefactional portion of the waveform in order to deliver the integrated energy in the

ultrasonic pulse accurately. An ideal hydrophone needs to possess:

- sufficient sensitivity for use in medical ultrasound applications

- uniform frequency response

- smallest possible active element in order to provide good directional

response and minimize spatial averaging effect

- good spatial resolution

- good signal fidelity

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- linearity in pressure level

- long-term stability in sensitivity

- known temperature dependence

Selection of piezofilm

Resonance frequency

In order to have a wider frequency response, thin film should be used. A thinner film

will have smaller active volume of piezoelectric material compared to thicker film

resulting in reduced sensitivity. The thickness mode resonant frequency (fr) is

determined by:

fr = t

c2

(A1)

where, c is the speed of sound in PVDF film and t is the total thickness of the films.

Electrical Impedance

The capacitance (C0) and the electrical impedance (Z) of the sensing element of the

hydrophone is given by:

C0 = tASε ; Z =

0

1Cjω

= Aj

tSωε

(A2)

where,

t is the thickness

A is the area of the active element

εS is the clamped dielectric constant

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It may be seen that the electrical impedance varies inversely proportional to the area of

the element size and directly proportional to the thickness of the PVDF film.

Electrode dimension

IEC 1102 specifies that the maximum effective radius amax should be [23]

amax = 1/22 0.25) (F4

+λ (A3)

where F is the focal number (ratio of separation distance between transducer and

hydrophone to the diameter of the probe). If the ultrasonic pulse contains harmonics, a

smaller spot size is required and the effective radius in such case should be [23]:

amax = F4λ n-1/2 (A4)

Taking into consideration of the above aspects, the design approach was formulated to

provide a rugged device capable of giving reasonably high sensitivity and broad

frequency response. The design involved laminating two layers of PVDF film in such a

way that the electrode and its leg are inside the lamination. Most of the outer side of

both the layers is metalized with gold and connected to the ground, which acts inherently

as a shielding to the live electrodes, thus reducing the noise pick up through radio

frequency interference. The laminated double layer structure has excellent noise

immunity.

The live electrode was chosen as a small circular dot having geometrical diameter of 0.4

or 0.6 mm, which was defined by vacuum deposited gold on top of chromium (chromium

was used for better adhesion) on one side of one layer. The specific central circular

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region (element spot) was poled so as to make it piezoelectrically active. The voltage

signal generated by the piezoelectrically active dot electrode was taken via a thin strip of

about 0.2 mm wide, gold deposited electrode leg, which extended to a point at the

periphery of the PVDF film and served as an electrical signal lead. The thickness of the

electrode and electrode-leg materials were chosen to be sufficiently thick to provide

durability and long term stability of the electrode but thin enough to avoid acoustic

loading of the PVDF film, thereby reducing the bandwidth. The size of the element leg

was also carefully chosen keeping in mind the effect of lead resistance.

The electrode was evaporated on one surface of the one layer, which is called active

layer. The outer surface of this active layer was divided into four quadrants and coated

with gold in such a way that a small portion of the electrode leg was visible, which

facilitated proper orientation of the hydrophone (the active element side of the

hydrophone facing the acoustic source) during measurement. The other layer, which is

called ground plane, was coated completely with gold. Although this protective coating

has some drawbacks, it was used in order to have a better signal fidelity by way of

achieving good grounding. It provides effective shielding to the live electrode from the

water in the test tank so that the capacitive loading of the element is reduced. Although

the coated double layer hydrophone reduces the electrical pickup, it has larger reflection

coefficient compared to single layer non-coated (coplanar) hydrophone.

Since the hydrophone has capacitive output impedance, it needs to be connected to an

amplifier having high input impedance in order to record the pressure waveform

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accurately. The preamplifier needs to be placed as close to the hydrophone as possible to

avoid signal loss in the cable. It maintains the voltage output of the hydrophone

unaltered over the frequency bandwidth of interest and it also acts as an impedance

matching network.

Although the element size was 0.4 or 0.6 mm in diameter, the overall size of the PVDF

film was designed to have 7 cm in diameter, which is fairly larger than the acoustic field

to be normally measured, so as to allow the beam from the imaging probe to pass through

freely thus facilitating acoustic transparency. The spot poling helped in preventing false

signals from acoustic perturbations and the absence of backing prevented reverberation

effects. Since the impedance of PVDF (4 MRayls) is close to that of water (1.5 MRayl),

the acoustic reflection coefficient at the surface of the PVDF film will be low for f < fr.

Moreover, since the transverse and radial mode of vibration are related to the dimension

of the entire hydrophone instead of the dimension of the active element alone, these

modes of vibrations are limited to low frequency level.

e. Fabrication Process

Electrode patterning was done using “shadow masking” procedure. In this method, a

metal mask covers the entire portion except the intended electrode pattern. A set of two-

piece sputtering mask was made from a thin stainless steel sheet by “Electro Discharge

Machine” (EDM). A thin plate was required for the mask to avoid shadowing of the

intended electrode area by the edge of the mask.

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A biaxially stretched and unpoled PVDF film was procured and kept between the two

plates of the mask and pinned and held firmly. The electrode and electrical leg was

vacuum deposited through metallic mask with a thin layer of chromium for better

adhesion and followed by slightly thicker layer of gold for good conduction. One side of

the 25 µm film was patterned for the active spot electrode and electrode leg and the other

side of the film was patterned with four quadrants for the ground plane. The second film

with 9 µm thick PVDF film was evaporated with gold on one surface, which is called the

second ground plane.

Both the ground layer and the active layer were separately stretched over supporting

rings. The protective gold coating improves the life and signal/noise level. In order to

have low contact resistance, gold plated surface mounting contacts were fixed directly on

to the film surface, making permanent electrical connection with the deposited electrode

leg by using conductive silver epoxy. Poling was performed on the clamped active layer

by heating it to a temperature well below its Curie/melt temperature while simultaneously

applying a high voltage for certain period of time, and then allowing it to cool down to

room temperature in the presence of the applied electrical field. The poling took place in

an air oven. The direction of the applied high voltage decides the direction in which the

piezoelectric effect is required. The poling field was applied perpendicular to the film

surface. This direction is defined as z coordinate, while the stretching direction is taken

as x, y axes. This polarization process rotates and aligns the dipole moments all in the

same direction, which is along the direction of polarization. After spot poling, the small

central region became piezoelectrically active.

159

The assembled and poled active layer and the unpoled ground layer were then laminated

together by applying a pressure with a dedicated press and keeping small drops of non-

conductive low viscosity epoxy between the two layers. It is called “drop method of

adhesion”. The laminating structure was kept pressed in the pressing mechanism until

the epoxy was cured at a set elevated temperature. The contacts of the electrode were

soldered to a standard 50 Ω coaxial cable. The outer sides (ground plane) of the

laminated films were connected to the screen of the coaxial cable, which was connected

to the ground, which minimizes the electrical radio frequency (RF) pickup. The

laminated films were mounted on to a hoop machined from ABS (Acrylonityile

Butadiene Styrene) material having an outer and inner diameter of 11 and 7 cm,

respectively and sealed effectively with pre-tapped mounting holes.

Since the voltage output from the active element was reduced due to the capacitance of

the cable, an unity gain preamplifier with high input impedance of 10 MΩ and 6 pF

having an output impedance of 50 Ω was fabricated in order to drive the electrode leg and

a short length (about 3 cm) of output cable. The preamplifier was assembled and potted

separately and connected close to the hydrophone using a removable connection. This

was done taking into account the manufacturing difficulty and practical usage, although

keeping the amplifier closest to the hydrophone is preferred to avoid signal loss in the

cable. If longer cable is used, the electrical reflections from each end of the cable will

distort the frequency response. Figure 40 shows the view of typical double layer

hydrophone developed.

160

f. Measurement Setup

Frequency Response Measurements

The details of the measurement setup used for the measurement of frequency response

and sensitivity were described in references 2 and 3.

Directivity Measurements

A broadband PZT (Lead Zirconate Ttitanate) transducer was mounted in a tank

containing deionized and degassed water at room temperature. The hydrophone under

test was placed facing the source at a known separation distance in the far field. It was

rigidly held by a mechanical holding device ensuring that the probe is positioned

perpendicular to the plane of the holder. The holding device enabled the hydrophone to

be rotated about an axis perpendicular to the beam alignment axis. The mounting

arrangements featured independent translation in the X, Y and Z directions besides

Figure 40: Actual view of the newly developed double layer polymer hydrophone probe

161

allowing the hydrophone to rotate its acoustical axis, and tilt and adjustment about the

vertical axis. The hydrophone movement was controlled through a stepping motor. The

spectrum analyzer and the stepper motor controller were interfaced to a personal

computer through an IEEE-488 bus for capture and storage of the hydrophone signals.

g. Measurement Procedure

Frequency Response Measurements

The frequency response and the sensitivity were measured using Time Delay

Spectrometry (TDS) combined with substitution method [3, 4]. The measurement

procedure is already outlined in the main text of the thesis.

For Directivity Measurements

The hydrophone was placed in water tank approximately 30 minutes before the

measurement began. The wideband source transducer was initially aligned visually,

directly facing the hydrophone such that the acoustic beam axis was aligned with the

hydrophone axis. The directivity measurements were performed by employing Time

Delay Spectrometry (TDS) technique [31, 52]. The hydrophone rotational geometry was

set up by using a special manipulator to rotate the hydrophone about its face until the face

of the hydrophone lay approximately perpendicular to the line of the transmitter acoustic

axis, and then rotating it slightly to maximize the output signal from the hydrophone,

which was displayed on the spectrum analyzer. This position was defined as zero degree

rotation. The hydrophone mount was rotated up to 500 on either side of its acoustical axis

and the hydrophone response was captured at every one-degree intervals. The

162

hydrophone was rotated such that the center of its active element was kept at the same

place in the ultrasound field. All measurements were performed in the far field of the

wideband acoustic source. The relative amplitude (dB) of the hydrophone response

measured as a function of frequency at a constant angle of rotation and for several values

of the angle of rotation was displayed on the spectrum analyzer. The displayed spectrum

of the signal represents the directivity patterns of the receiver as a continuous function of

frequency. The spectral data from the hydrophone measurements were transferred from

the spectrum analyzer to the controlling computer. From the spectra corresponding to

different incident angles, the frequencies corresponding to –3 dB and –6 dB points were

found. The data obtained was rearranged to yield the hydrophone output voltage as a

function of the angle of rotation at a constant frequency of interest and analyzed by

displaying all the angular data relative to 00 at the specific frequency. The angles to the

–3 dB and –6 dB points were used to find the effective diameter of the hydrophone. The

measured 3 dB and 6 dB half angles of the hydrophone response at a particular frequency

gave the effective radius. If the effective radius a3 and a6 are equal to each other within

±10% of the maximum value, then the mean value shall be used as the effective radius.

If not, the value of the two whose corresponding angle θ is closer to 100 shall be used

[61]. The mean of those two values resulted the effective radius of the hydrophone at

that frequency. If θ6dB is larger than 300, then θ3dB was used to find the effective radius.

In this study, the angular response of the hydrophones was analyzed at the frequency of 5

MHz, 7.5 MHz and 10 MHz. The effective radii determined at different frequencies (e.g.

5 MHz, 7.5 MHz and 10 MHz) were averaged in order to determine the mean effective

diameter.

163

h. Theory and Experiments

Acoustic Sensitivity

The sensitivity of a hydrophone depends on the thickness resonance of the film and the

electric and piezoelectric properties of the polymer material used. The sensitivity is

determined by the active area of the element and the capacitance loading of the spot

electrode, electrode leg and the connecting cable over a frequency range. Since the

thickness resonance is inversely proportional to the thickness of the PVDF film, there is

an initial gradual increase in sensitivity with frequency. At the same time, the electrode

capacitance decreases with increasing frequency due to decrease in the dielectric constant

of PVDF material, causing significant electrical loading of the active element. Hence the

sensitivity of the hydrophone increases slowly with increasing frequency [49].

If the acoustic wave having a frequency below the resonance is normally incident on the

active element, the proportionality of the output voltage of the hydrophone is given by:

V ∝ ce

e

CCgtpC+

(A5)

where,

p is the acoustic pressure incident on the active element

g is the receiving constant of the PVDF film

Ce is the element capacitance

Cc is the cable capacitance, including the shielding film capacitance

t is thickness of the active layer

The schematic representation showing the capacitance loading of the hydrophone is given

in Figure 41.

164

HydrophoneSensing Element

Coaxial Cable

Ce

VocCs Cc CL

Moc Mec Meff

Measuring Instrument(Spectrum Analyzer/Oscilloscope)

where,

Ce is the element capacitance

Cs is the stray capacitance


CL is the load capacitance (measuring instrument)

MOC is the open circuit sensitivity

MEOC is the end of cable sensitivity

Meff is the effective sensitivity

As shown in the Figure 41, the element capacitance Ce is loaded by the stray capacitance

Cs and the cable capacitance Cc, including the shielding film capacitance. In order to

obtain broader bandwidth, thinner films are used. Thinner the film, the element

capacitance is more, resulting in lower impedance.

Figure 41: Schematic representation showing the capacitance loading of the hydrophone probe

165

The active film element and the shielding film capacitances and the active film

impedance are given, respectively by:

Ce = tAsε

, Cp = p

s

tAε

, Z = Aj

tSωε

(A6)

where

εS is the clamped dielectric constant of PVDF

t and tp are thicknesses of the active and shielding films, respectively

A is the area of the electrode

Better resolution/directionality can be obtained by using a small element area and the

electrical impedance is increased as the element diameter decreases.

The end of cable open circuit sensitivity (MEOC) of the spot poled hydrophone is given by

the ratio of the voltage developed and the acoustic pressure incident on the active

element, which can be theoretically calculated by using the following expression [64]:

MEOC = pV =

ce

e

CCgtC+

(A7)

where,

V is the end of cable open circuit voltage

p is the acoustic pressure in Pascal incident on the hydrophone

t is the thickness of the active layer sensing element

g is the piezoelectric receiving constant of the PVDF material

Ce is the capacitance of the active element of the hydrophone


166

The capacitance of the element was measured at low frequency far below the resonance

(i.e. 1 kHz) by using capacitance meter. Since the open circuit voltage sensitivity (MOC)

(which is the actual output of the hydrophone) is difficult to measure, it can be calculated

from the measured end of cable sensitivity (MEOC) of the hydrophone, by measuring the

capacitance of the element using an impedance analyzer. For the hydrophone having no

preamplifier, it is only a capacitive load and the relationship is given by:

MEOC = MOC

++ cse

e

CCCC

(A8)

The effective sensitivity (Meff ) is given by:

Meff = MOC

+++ Lcse

e

CCCCC

(A9)

It may be seen that the capacitance of the cable greatly influences the voltage sensitivity

of the hydrophone. To alleviate this problem, a preamplifier was placed as close to the

hydrophone output terminal as possible.

While using the hydrophone with preamplifier, the relationship is given by:

Meff = MOC

+++

+22

22

)]()([)]()([)()(

ZIZIZRZRZIZR

mLmeLe

LmLe (A10)

where,

Z is the measured complex impedance of the hydrophone

ZL is the input impedance of the measurement device (e.g. Oscilloscope)

Re and Im denote the real and imaginary parts of the complex impedance.

167

The preamplifier of the hydrophone usually has an output impedance of 50 Ω to drive the

50 Ω cable. If the input impedance of the measuring instrument has 50 Ω load, it

matches with the output impedance of the preamplifier of the hydrophone. Under such

condition, the end of cable sensitivity can be recorded as it is without making any

corrections using the expression given in A10. It should always be ensured that the

electrical loading condition at the time of calibration is the same as during the time of use

so that there will not be any change in sensitivity. The evaluation of the performance

characteristics of several hydrophones revealed that larger diameter hydrophones

exhibited better sensitivity. It was also found that sensitivity varies from one hydrophone

to another within the same type, due to variability in the fabrication process, contributing

different capacitive load and also due to the variation of effective aperture size due to

fringe field effects during electrical poling.

Table 2. Comparative statement showing the variation of frequency response and sensitivity of hydrophones having different thickness of PVDF film. Type of Construction Theoretical Typical Typical

Thickness Measured Sensitivity

Resonance Frequency dB re 1V/µP

Frequency Response

Hydrophone using:

25+25 µm thick PVDF film

(0.4 mm spot size) (with preamp) 22 MHz 1 - 20 MHz -252





168

From the details provided in the table above, it is seen that increase in sensitivity has

been demonstrated using 25+9 µm thick PVDF film compared to 9+9 µm thick PVDF

film.

Frequency Response and Bandwidth

The resonance frequency is inversely proportional to the thickness of the PVDF film and

the frequency response is relatively constant up to its resonance frequency. Different

configuration of double-layer hydrophone probes made by employing 25 and 9 µm thick

PVDF films having 0.4 and 0.6 mm of geometrical spot diameter were tested for their

performance. The frequency responses of the hydrophones were investigated using two

types of measurements: (i) one was the absolute calibration of randomly selected

hydrophone at National Physical Laboratory, UK and (ii) the second involved comparing

the frequency response of the newly developed hydrophone with a reference hydrophone

which was already calibrated at NPL. The results of the randomly chosen hydrophone

probes were also verified at National Physical Laboratory (NPL), UK. The frequency

response of a double-layer hydrophone probe fabricated with (9+9) µm thick PVDF film

measured at NPL showed that the sensitivity is being constant to ±3 dB between 1 to 50

MHz is shown in Figure 42, which is an example of wider bandwidth using a thinner

PVDF film.

169

In the case of hydrophone probe fabricated with 25+25 µm thick PVDF film, the

sensitivity variation in the frequency range from 1 to 20 MHz is within ±3 dB of the

mean value for all frequencies. The polymer hydrophones developed using dissimilar

thickness (25+9 µm) have demonstrated higher bandwidth compared to that of

hydrophones made using similar thickness (25+25 µm) of PVDF film as seen from the

data provided in Table 2. A flat frequency response up to 25 MHz has been

demonstrated as shown in Figure 43.

The downshift of the resonance frequency of hydrophone using 25+9 µm thick PVDF

film from the theoretical calculation of 32 MHz was due to the thickness of bonding

layer, damping effect of the electrode material and the attenuation in the PVDF.

Figure 42: Frequency response plot of a double layer hydrophone probe using 9+9 µm thick polymer film.

-280

-260

-240

-220

-200

0 10 20 30 40 50

Frequency in MHz

End

-of-

Cab

le S

ensi

tivity

(dB

re

1V/u

Pa)

9+9 micron thick PVDF film

170

Hydrophone Effective Diameter

Theory

If the effective diameter of the hydrophone is large compared to the acoustic wavelength

and ultrasound beam dimension, the measured acoustic pressure amplitude will be less

than the actual value. This is because the hydrophone will respond to the pressure

averaged over its active element and hence errors in spatial-peak pressure measurement

will occur. This can lead to underestimates of the true acoustic pressures and derived

intensities, an effect that has been termed as spatial averaging [50]. Spatial distortion of

the beam profile will result if the effective diameter of the hydrophone is large in

comparison with the beam dimension. It will be under sampled, and the information

carried in spatial harmonics above a certain frequency would be lost. A correction factor

-280

-260

-240

-220

-200


Vol

tage

sens

itivi

ty (d

B r

e 1V

/uPa

)




Figure 43: Frequency response plots of double layer hydrophones using 25+25 µm and 25+9 µm polymer film

171

has to be used to compensate these errors in the measurements data. As per the

AIUM/NEMA [65], the guideline for choosing the effective hydrophone diameter, de is:

de < ds2zλ if z/ds ≥ 1 and de <

2λ if z/ds < 1 (A11)

where

ds is the source diameter

z is the distance from the hydrophone to the source

λ is the acoustic wavelength.

The hydrophone diameter is found based on a comparison of the hydrophone's directional

response with theoretical directional response of a uniform receiver with circular aperture

[65]. Hydrophone's response in the form of the amplitude of the received waveform is

plotted as the function of the angle of rotation θ, with respect to the geometrical axis of

the hydrophone. This angle is related to the effective diameter [66].

The directional response of the hydrophone was modeled taking into account the film

thickness and Lamb wave propagation. It is approximated to the first order Bessel

function of a stiff disc receiver. For circular hydrophone apertures of radius a, incident

waves of wavelength λ, and measured half angle θ, the far field pressure directivity

function is given by [67]

P(r,θ) = p(r)

θθ

sin)sin(2 1

kakaJ (A12)

The hydrophone voltage response in the receive mode is [67]

V (θ) = V (0)

θθ

sin)sin(2

1

kakaJ (A13)

172

where

V (0) is the on axis output voltage

a is the radius of the hydrophone

r is the distance from the center of the source to the point of measurement

θ is the angle between r and the source axis

J1 is the first order Bessel function.

k is the acoustic wave number (cω )

The directivity of an ideal stiff disc hydrophone of radius a is given by [64]:

D (θ, z) = θθ

ππ

sin)sin(2)( 1

2

kakaJ

zaj

× (A14)

where z is the axial distance from the source transducer. In practical reality, the

directivity of the hydrophone will deviate from the ideal condition.

The effective radius ae in mm of the active element of the hydrophone is determined from

the measured directional response by inserting the measured half angle θ of the directivity

function at the 3 dB and 6 dB points into the theoretical directivity function and

calculating the two values of the effective radius at a particular frequency employing the

following relation [61].

a3dB = dB3sinf2

c62.1θπ

and a6dB = θπ dB6sinf2c22.2

(A15)

where

f is the frequency of the angular measurements

c is the speed of sound in water at the particular temperature.

173

θ3dB, θ6dB are the measured half angle at which the hydrophone response is 3 dB

and 6 dB down from its reference level at θ = 00 incidence.

The two measured half angles (θ3 and θ6) at one frequency yield two theoretical values of

the hydrophone radius (a3 and a6), and the mean of those two values yields the effective

radius at that frequency [65].

Experimental Results

The directivity measurements were used to determine the effective size of the sensitivity

element of the hydrophone. It was seen that the effective diameter of the hydrophone

varied significantly with frequency and with reference to axis of rotation. It is essential

that the effective diameter be known as a function of frequency at specified interval in

order to apply the spatial averaging corrections [50, 62] while reporting the acoustic

measurement results on medical ultrasonic transducers.

Tables 3 and 4 contain data, which illustrate the effective diameters, measured along two

orthogonal axes on different kind, geometrical spot size, and having different thickness of

PVDF film hydrophones. The tables show the values of effective diameter of the

hydrophones d (0) and d (90) for two scan orientation, 00 and 900. The d (90) was the

result of directional response measurements conducted with the hydrophone being

collinear with the electrical leads. Normalization with respect to the hydrophone’s

geometrical diameter was accomplished for easy comparison. From the results, it was

seen that the effective diameter decreased with increasing frequency.

174

Table 3. Directivity data illustrating the variation of effective diameter of the hydrophones measured along the two orthogonal axes at 5, 7.5 and 10 MHz. SLID Active At 5 MHz At 7.5 MHz At 10 MHz Diameter Measured Diameter in mm Measured Diameter in mm Measured Diameter in mm in mm Normalized

Value

Normalized Value

NormalizedValue

d (0) d (90)

d (0) d (90)

d (0) d (90)

1 A1 0.4 0.507 0.802 1.582 0.492 0.779 1.583 0.491 0.775 1.578

2 A2 0.4 0.545 0.609 1.117 0.518 0.598 1.154 0.510 0.589 1.155

3 A3 0.4 0.540 0.636 1.178 0.517 0.617 1.193 0.505 0.603 1.194

4 A4 0.4 0.535 0.619 1.157 0.519 0.602 1.160 0.507 0.592 1.168

5 A5 0.4 0.532 0.621 1.167 0.512 0.605 1.182 0.489 0.599 1.225

6 A6 0.4 0.527 0.621 1.178 0.505 0.589 1.166 0.493 0.586 1.189

7 B1 0.5 0.544 0.563 1.035 0.514 0.540 1.051 0.508 0.541 1.065

8 C1 0.5 0.584 0.627 1.074 0.560 0.591 1.055 0.553 0.583 1.054

Table 4. Directivity data illustrating the variation of mean effective diameter of the hydrophones measured along the two orthogonal axes. SL Identification Active Mean value among 5 MHz, 7.5 MHz & 10 MHz Normalized Value Diameter

in mm da

Measured Diameter in mm d (0) d (90)

1 A1 0.4 0.497 0.785 1.243 1.963

2 A2 0.4 0.524 0.599 1.310 1.498

3 A3 0.4 0.521 0.619 1.303 1.548

4 A4 0.4 0.520 0.602 1.300 1.505

5 A5 0.4 0.511 0.608 1.278 1.520

6 A6 0.4 0.509 0.599 1.273 1.498

7 B1 0.5 0.522 0.548 1.044 1.096

8 C1 0.5 0.566 0.600 1.132 1.200

d(0))90(d

d(0))90(d

d(0))90(d

da)90(d

da)0(d

175

Typical directivity pattern of a hydrophone having 0.4 mm, 0.6 mm and 1 mm

geometrical spot sizes are presented in Figure 44 for comparison.

Directional Response

(a) (b)

(c)

Figure 44: Typical directional response of double layer polymer membrane hydrophones with active element size of 0.4 mm (Fig. a), 0.6 mm (Fig. b) and 1 mm (Fig. c) in diameter, measured at the frequencies of 5 MHz, 7.5 MHz and 10 MHz.

176

Directional Response

The membrane hydrophone could be susceptible to Lamb wave propagation across the

PVDF film surface (causing a build up of acoustic pressure at the active element)

resulting in side lobes when the sound is incident at the critical angle of ≈500 [68]. The

directional responses of membrane hydrophones are affected by the variation of

piezoelectric sensitivity of PVDF with the direction of acoustic stress and by Lamb wave

propagation in the membrane, which causes a build up of acoustic pressure at certain

angles of incidence. Both these effects cause the directivity of a hydrophone to deviate

from that of a stiff plane disc [69].

The angular responses of several double-layer PVDF hydrophones of different thickness

of PVDF film and various geometrical spot sizes were obtained by rotating the

hydrophone in the far field of the transducer’s plane wave and measuring the

hydrophone’s response at some angle of rotation. The plots are shown in Figure 45 and

46. It was seen that the Lamb wave propagation appears on the directional response plots

in the form of side lobes. It was also seen how the presence of Lamb wave affects the

main lobe at higher frequencies and also in larger diameter hydrophones. At lower

frequencies, the side lobes dwarf the main lobe. As the frequency increases, the main

lobe strengthens and becomes narrower in angular width as expected. The main lobe

gradually scaled down in amplitude at higher frequencies. It also revealed that the peaks

of the directional response curves at different frequencies are collinear. This means that

one could minimize the error while making measurements if the angular alignment of

hydrophone is optimized at the highest frequency.

177

0.4 mm dia spot size 0.6 mm dia spot size

1mm dia spot size

Figure 45: Combined directional response of double layer polymer membrane hydrophones with active element size of 0.4 mm in diameter, measured at the frequencies of 5 MHz, 7.5 MHz and 10 MHz.

Figure 46: Combined directional response of double layer polymer membrane hydrophones with active element size of 0.4 mm, 0.6 mm and 1 mm in diameter, measured at 7.5 MHz.

1 mm dia spot size

5 MHz 7.5 MHz

10 MHz

178

Orientation effects

To study the apodized behavior and asymmetry of the sensitive region of the hydrophone

probes, angular response of several hydrophone probes of different kinds, various

geometric spot size and different thickness of PVDF film were measured in two

rotational axes perpendicular to each other. The effective diameter of the hydrophone

probe was determined from the measurement of its directivity pattern.

Table 5. Directivity data illustrating the variation of effective diameter of the hydrophones having different thickness of PVDF film measured along the two orthogonal axes.

Type Membrane Geometrical Effective Spot Size % variation Thickness Spot Size d (0) d (90) d (0) d (90) Bilaminar 25+25 µm 0.4 mm 0.497 0.785 24% 96% Bilaminar 25+9 µm 0.4 mm 0.521 0.619 30% 55% Bilaminar 9+9 µm 0.4 mm 0.511 0.608 28% 52%

Sensitive region (0.4mm diameter)

Gold deposited electrical lead (0.2mm width)Hydrophone Orientation

d (0) d (90)Figure 47: Combined directional response of double layer polymer membrane hydrophone with active element size of 0.5 mm in diameter, measured at 10 MHz in two orthogonal axes.

179

From Figure 47 and Table 5, it may also be seen that there was a marked asymmetry in

the effective diameter measured along the two orthogonal axes. The diameter measured

along the line collinear with the electrical lead was greater of the other diameter. This

asymmetry of the sensitivity region of the membrane hydrophone was due to the fact that

the electrode contact region has distorted the spot size, which can be due to fringe field

effects during electrical poling of the PVDF film [68]. During the electrical spot poling,

the applied high electric field fringes beyond the edge of the spot electrode, and poled the

areas of the piezoelectric polymer beyond the intended spot diameter, thus increasing the

effective spot size. The distortion of the fringe field at the point where the electrode meet

the spot size is creating the asymmetry in the effective diameter which is true as

evidenced from the measurements conducted along the line collinear with the electrode.

The apodized behavior can be due to directional dependence of the piezoelectric

sensitivity and fringe poling field, and this fringe field also can lead to asymmetry of the

sensitive region [68]. For the larger spot size, the effect of the fringing field is

comparatively smaller than the hydrophone with smaller spot size.

Most of the hydrophones used in this study do not satisfy the AIUM/NEMA’s guideline

namely |a

aae − | < 0.3, while the measurements were made by keeping the hydrophone

along the line collinear with the electrical leads.

There will also be a variation of measured voltage sensitivity depending upon which side

of the hydrophone is facing the source transducer [49].

180

Long Term Stability

The initial absolute calibration of a newly developed hydrophone was undertaken at NPL

and it was recalibrated at the same facility (NPL) every year after routine usage in

measuring the acoustic output of medical imaging transducers on most of the days in

every week throughout the year. No change in sensitivity was observed over a two-year

period of time, which shows the temporal stability of the hydrophone in routine use.

i. Summary

Several double layer hydrophone probes using 25+25 µm and 9+9 µm thick PVDF films

and characterized for

- acoustic sensitivity

- frequency response

- bandwidth

- angular response

- effective aperture size

- orientation effects

and the results agree with the theory and demonstrated good performance. The improved

design of double layer hydrophone probe using dissimilar thickness (25+9 µm) of PVDF

film developed in this work exhibited wider bandwidth compared to the hydrophone

using 25+25 µm thick PVDF film better sensitivity and robust compared to the

hydrophone using 9+9 µm thick PVDF film. The hydrophones using 9+9 µm film were

in fact much more prone to suffer damage than 25 µm thick film.

181

j. Conclusions

The results of this research concludes the following:

(a) This improved design of double-layer hydrophone probe to enhance the bandwidth

using dissimilar thickness of (25+9 µm) PVDF film developed in this work exhibited

higher bandwidth compared to 25+25 µm thick film and possess better sensitivity and

robustness compared to 9+9 µm thick PVDF film as evident from the data provided in

Table 2. This was the first one of this kind and no one has developed this type of

dissimilar thickness hydrophone previously.

(b) The directionality was found affected by the angular variation of piezoelectric

sensitivity and also for large angle of incidence by Lamb wave propagation in the

membrane.

(c) The effective diameter of the hydrophone varied significantly with frequency and with

reference to axis of rotation. Since the effective diameter values obtained from the

measurements at two perpendicular axes were more than ±10% of the maximum value

for most of the hydrophones, it emphasized the necessity that the amplitude directivity

does need to be measured at different frequencies and at two perpendicular axes and the

values of both needs to be provided. The relevance of this effect in practice is that the

orientation of the membrane hydrophone must always be the same, whether it is being

calibrated or being used as reference device.

182

(d) While reporting measurement results on medical ultrasound equipments, one needs to

apply spatial averaging corrections to the results obtained at a particular operating

frequency of the transducer. In view of the usage of transducers with enhanced upper

bound of the operating frequencies (both fundamental and harmonics) for the diagnostic

imaging at present, it is considered necessary that the effective diameter of the

hydrophone be known as a function of frequencies at specified intervals to cover the

whole frequency bandwidth of the hydrophone including the upper limit of the frequency

bandwidth.

183

VITA

Vadivel Devaraju received his Bachelor of Engineering (B.E) degree in Electrical Engineering from the University of Madras, India, and Master of Science (MS) in Biomedical Engineering from Drexel University, Philadelphia, Pennsylvania in 1995. His masters’ thesis focused on calibration of ultrasonic hydrophones using Time Delay Spectrometry technique. While pursuing his studies in biomedical engineering, he worked at the medical imaging division of the department of radiology at the University of Pennsylvania for about 2 years. He also worked as the manager of ultrasound measurement laboratory at Preceptron/Sonic Technology for about 3 years, where his primary responsibilities involved: • Testing, characterizing, measuring and evaluating a wide variety of medical ultrasound

equipment and other medical devices from various manufacturers worldwide to verify compliance with US FDA 510(k) regulations and IEC standards. The ultrasound systems included A-mode, M-mode, B-mode and pw/cw/color/power Doppler mode and in addition to fetal heart monitors, cardiac output monitoring system, ultrasonic surgical devices, ultrasonic physiotherapy equipment, active implantable medical devices, high frequency intravascular devices etc. He has tested ultrasound devices from 20kHz to 50MHz, over a wide range of power levels and prepared 510(k) reports for the above-mentioned systems for submission to FDA.

• He has manufactured bilaminar PVDF membrane hydrophones featuring wide bandwidth and a selection of active element sizes and exported them to medical ultrasound equipment manufacturers, research organizations and universities throughout the world. His employer was the only domestic manufacturer of such device and one of the two manufacturers in the world.

• He has calibrated the newly built hydrophones, and performed the directivity measurements to ascertain the active element size.

• He has manufactured and tested self-monitoring shockwave hydrophone system, suitable for quantitative measurements of lithotripters and other shockwave sources.

The highlights of his PhD work included: • Development and characterization of double layer hydrophone probes using dissimilar

thickness of PVDF film, which demonstrated simultaneous enhancement of bandwidth and sensitivity.

• Development of a new calibration method using planar scanning technique to calibrate the hydrophone probe below 1 MHz in order to address the issue of non-existence of calibration standard.

• Development of single element, single layer transducer and multilayer Barker coded polymer transducers and characterization in pulse echo mode, which exhibited widest bandwidth and increased pulse echo sensitivity.

• Design, development and characterization of ultra-wideband single layer array transducer and multilayer Barker coded polymer array transducers for medical ultrasound imaging applications, which are capable of operating at all clinically relevant frequencies using single scanhead and suitable for tissue and contrast specific harmonics imaging.

His published work relating to this thesis has been cited under bibliography.

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