bioimpedance mapping of the cervix · 2010-06-09 · bioimpedance spectroscopy has shown potential...

QUEENSLAND UNIVERSITY OF TECHNOLOGY

SCHOOL OF PHYSICAL AND CHEMICAL SCIENCES

Bioimpedance Mapping

of the Cervix

Submitted by Jye Geoffrey SMITH to the School of Physical and Chemical Sciences, Queensland University of Technology, in partial fulfilment of the requirements of the degree of Doctor of Philosophy.

May 2008

Keywords

Bioimpedance, Mulitifrequency, Impedance, Cervical Cancer, Bovine Blood,

Haematocrit, Tetrapolar, Finite Element Analysis, Impedance Mapping, Mul-

tiplexer

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Abstract

Bioimpedance spectroscopy has shown potential as a method for characterising

biological tissue with the use of a tetrapolar electrode configuration. Brown

et al. (2000) demonstrated that the configuration is capable of distinguishing

between normal squamous epithelium and Cervical Intra-epithelial Neoplasia

(CIN). However little has been done to identify the volumes of tissue that

contribute to the measured impedance. Brown et al. employed a probe with

a single tetrapolar electrode set thus analysing single points of tissue. The

probe was required to be moved in order to “sample” other areas of tissue.

This method provides no spatial information of the lesion boundaries.

The overall objective of this research was to design and construct an impedance

mapping system (IMS) for objective virtual biopsy of lesions by bioimpedance

spectroscopy (BIS). Initially freshly excised cervical tissue was to be tested

however as the study progressed this proved problematic and bovine blood

was chosen as a suitable substitute.

Specific aims were to; Investigate the spatial sensitivity distribution of the tetrapolar electrode

configuration via finite element analysis (FEA).

v

Design a novel front end multiplexing system and multi-electrode array

for mapping the impedance of the tissue of interest. Experimentally confirm the efficacy of the approach to identify regions of

different impedances and their boundaries using bioimpedance mapping.

The present study used finite element analysis (FEA) to investigate the

spatial variation in sensitivity of the tetrapolar electrode configuration and

identify which volumes of tissue were included in the measured impedance.

An impedance mapping device was also designed and constructed utilising

the tetrapolar electrode configuration in an expanded array of 25 electrodes.

This array allowed the surface of an area of tissue to be mapped and lesion

boundaries identified in an objective manner.

FEA was also used to model lesions in healthy tissue and the sensitivity

fields associated with the tetrapolar configuration. The FEA indicated that

anomalous results would be obtained when a lesion was located between a

drive and measurement electrode pair. In this case the lesion resulted in an

increase in impedance with respect to the impedance of healthy tissue, whereas

a lesion should result in a decrease in measured impedance relative to that of

healthy tissue. The anomaly was found to produce false negative results for

small lesions up to 0.4 mm and even a lesion with radius of approximately

0.75 mm could be undetected as the measured impedance spectrum for such

a lesion is similar to that of healthy tissue. Modelling also provided insight

into the sensitivity fields for an electrode array and its efficacy in accurately

measuring the surface impedance of tissue and lesions of interest.

The impedance mapping system (IMS) developed used an array of 25 (5x5)

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electrodes. The array allows 64 individual tetrapolar measurements to be

obtained at 16 locations, providing an impedance map of 49 mm2 on the surface

of a tissue sample. Multiple measurements at each location reduce the chance

of anomalous results since these can be identified and excluded. Software was

developed to display the measured impedance maps and regions of different

impedance were easily identified

Testing of the IMS using bovine blood showed separation of the measured

impedance for a range of haematocrit between 0 - 80%. Introduced volumes of

red blood cells (RBC) or clots (to mimic lesions) to the plasma (haematocrit

0%) were also clearly identified using the IMS. It was seen that measurements

made at the boundary of 2 different haematocrits (ie 2 volumes of different

impedance) resulted in an anomalous result as indicated by the FEA modelling.

However it was demonstrated that these anomalies can be used to objectively

identify the introduced RBC (lesion) boundaries.

A more efficient electrode stepping sequence was also developed taking

advantage of the reciprocal nature of the tetrapolar electrode configuration.

This development allows for the electrode array to be doubled in size using

the same components, and to sample twice the surface area in the same time

taken using the initially developed system.

In summary, an impedance mapping system has been modelled, designed

and developed for tissue characterisation by bioimpedance measurements. The

technique has been shown experimentally to be able to detect regions of differ-

ent impedance and is in agreement with the finite element analysis performed.

Further development of the IMS will allow progressive monitoring of suspect

lesions in-vivo and better identification of their spatial distribution for biopsy.

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List of publications

1. Smith, J. G., Thomas, B. J. & Cornish, B. H. (2004). FEA modelling: ef-

ficacy of tetrapolar electrode arrays in virtual biopsy-cervix, XII ICEBI.

2. Smith, J. G., Thomas, B. J. & Cornish, B. H. (2007). A Pilot Study For

Tissue Characterisation Using Bioimpedance Mapping, XII ICEBI.

3. Smith, J. G., Thomas, B. J. & Cornish, B. H. (2007). Queensland Univer-

sity of Technology. Impedance Measurement Process. Australian Patent

Application Number: 2007904287.

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Contents

1 Introduction 1

1.1 Cervical Cancer . . . . . . . . . . . . . . . . . . . . . . . . . . . 1

1.2 Bioelectrical Impedance Analysis . . . . . . . . . . . . . . . . . 3

1.2.1 Bioimpedance Background . . . . . . . . . . . . . . . . . 3

1.2.2 Biological Tissue Equivalent Circuit . . . . . . . . . . . . 5

1.2.3 Applications of BIA to Tissue Characterisation . . . . . 9

1.2.4 Tetrapolar Electrode Configuration . . . . . . . . . . . . 12

1.3 Finite Element Analysis (FEA) . . . . . . . . . . . . . . . . . . 13

1.4 Aims and Objectives . . . . . . . . . . . . . . . . . . . . . . . . 15

2 Modelling of Sensitivity Distributions 17

2.1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17

2.2 Lesion Modelling . . . . . . . . . . . . . . . . . . . . . . . . . . 18

2.2.1 Methodology . . . . . . . . . . . . . . . . . . . . . . . . 18

2.2.2 Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22

2.3 Sensitivity Fields . . . . . . . . . . . . . . . . . . . . . . . . . . 27

2.3.1 3-D Sensitivity Field Model . . . . . . . . . . . . . . . . 28

2.3.2 Electrode Array Sensitivity Fields . . . . . . . . . . . . . 36

3 Impedance Mapping System Overview, Operation and Test-

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ing 43

3.1 Current Source & Potential Measurement . . . . . . . . . . . . . 45

3.2 Multiplexers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 45

3.3 Electrode Array . . . . . . . . . . . . . . . . . . . . . . . . . . . 46

3.4 Impedance Data Analysis . . . . . . . . . . . . . . . . . . . . . 49

3.5 Impedance Mapping System Graphical Use Interface and Oper-

ation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 50

3.6 Impedance Mapping System Testing . . . . . . . . . . . . . . . 51

3.6.1 Testing of Multiplexer Ron Contribution . . . . . . . . . 51

3.6.2 RRC Circuits . . . . . . . . . . . . . . . . . . . . . . . . 56

3.6.3 Resistor Matrix . . . . . . . . . . . . . . . . . . . . . . . 59

3.6.4 Biological Tissue . . . . . . . . . . . . . . . . . . . . . . 62

3.6.5 Testing Summary . . . . . . . . . . . . . . . . . . . . . . 64

4 Bioimpedance Mapping - Results and Discussion 67

4.1 Effect of Tissue Sample Size . . . . . . . . . . . . . . . . . . . . 68

4.1.1 Electrode Array Coverage . . . . . . . . . . . . . . . . . 68

4.1.2 The Effect of Sample Thickness . . . . . . . . . . . . . . 71

4.2 Homogeneous Haematocrit Impedance Mapping . . . . . . . . . 73

4.3 Impedance Maps . . . . . . . . . . . . . . . . . . . . . . . . . . 76

4.3.1 Plasma with Introduced Red Blood Cells . . . . . . . . . 76

4.3.2 Plasma with Introduced Red Blood Cell Clot . . . . . . 80

4.4 Experimental Comparison with Modelled Sensitivity Field . . . 85

4.4.1 Anomalous Measurements . . . . . . . . . . . . . . . . . 85

4.4.2 Reciprocal Electrodes . . . . . . . . . . . . . . . . . . . . 87

4.5 Lesion Boundary Identification . . . . . . . . . . . . . . . . . . . 92

4.6 Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 95

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5 Conclusion 97

Bibliography 101

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List of Figures

1.1 Schematic of modelled tissue when constructed of its basic com-

ponents. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

1.2 (a) Current pathways at low frequencies. (b) Current pathways

at high frequencies. . . . . . . . . . . . . . . . . . . . . . . . . . 6

1.3 Impedance spectrum for modelled tissue when constructed of

its basic components. . . . . . . . . . . . . . . . . . . . . . . . . 7

1.4 Impedance spectrum for biological tissue with a depressed semi-

circle. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

1.5 Changes in tissue structure associated with the progression of

CIN in cervical squamous epithelium (Walker et al., 2000). . . 10

1.6 Tetrapolar arrangement used for taking of bioelectrical impedance

measurements. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12

1.7 (a) Cross section of finite element model of single cell. (b) Finite

element tissue model (Walker et al., 2000, 2001a, 2001b) . . . . 14

2.1 (a) Tetrapolar electrode model with the lesion located between

a drive and measurement electrode. (b) 3D view showing the

lesion located immediately below the surface of healthy tissue. . 20

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2.2 Experimental electrode placement on the surface of a saline so-

lution and brass rod with radius ’r’ located (a) central to the

electrodes (b) between a drive and measurement electrode. . . . 21

2.3 Modelled results of tetrapolar electrode configuration for a le-

sion located centrally in the electrode configuration. . . . . . . . 24


sion located midway between a drive and measurement electrode

(Legend displays frequency in kHz). . . . . . . . . . . . . . . . . 25


sion located midway between a drive and measurement electrode

(Legend displays lesion radius in mm). . . . . . . . . . . . . . . 25

2.6 Measurement with (a) no lesion and (b) a lesion of radius 0.75

mm in otherwise healthy tissue. . . . . . . . . . . . . . . . . . . 26

2.7 Experimental and modelled results of the tetrapolar electrode

configuration for a brass rod located centrally in the electrode

configuration. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 26

2.8 Experimental and modelled results of the tetrapolar electrode

configuration for a brass rod located midway between a drive

and measurement electrode. . . . . . . . . . . . . . . . . . . . . 27

2.9 3-D modelled tetrapolar configuration and current density con-

fined by the boundaries. . . . . . . . . . . . . . . . . . . . . . . 30

2.10 Sensitivity field at a depth of 0.0 mm. . . . . . . . . . . . . . . . 30





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2.16 Maximum sensitivity in each plane as a function of depth for

electrode spacing’s 0.2 mm, 0.4 mm, 0.6 mm, 0.8 mm and 1.0

mm. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 34

2.17 Sum of the absolute sensitivity in each plane as a function of

depth for electrode spacing’s 0.2 mm, 0.4 mm, 0.6 mm, 0.8 mm

and 1.0 mm. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 35

2.18 Measured potential against tissue medium thickness for differing

electrode spacings. . . . . . . . . . . . . . . . . . . . . . . . . . 36

2.19 The four unique electrode sites modelled. . . . . . . . . . . . . . 37

2.20 Maximum sensitivity in each plane as a function of depth for

electrode sites 1, 2, 3 and 4 (see figure 2.19). . . . . . . . . . . . 38

2.21 Sum of the absolute sensitivity in each plane as a function of

depth for electrode sites 1, 2, 3 and 4(see figure 2.19). Locations

1 and 3 are difficult to see due to overlap. . . . . . . . . . . . . 39

2.22 Sensitivity field at a depth of 0.1 mm for electrode site 1 shown

in figure 2.19. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 39

2.23 Measurement thickness for an array with 0.77 mm electrode

spacing and a single electrode configuration with the same spacing. 40

2.24 Sensitivity field at a depth of 0.1 mm for a tetrapolar electrode

configuration with inactive electrode inside the configuration. . . 41

3.1 Schematic of the bioimpedance mapping system. . . . . . . . . . 44

3.2 Functional block diagram and pin configuration of multiplexer

ADG732. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 46

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3.3 Multiplexer circuit diagram. Parallel port represented by Header

9 and electrode array socket by Header 25. SFB7 drive and mea-

surement electrodes connect through red, black and white, blue

respectively. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47

3.4 PCB electrode array. . . . . . . . . . . . . . . . . . . . . . . . . 48

3.5 Electrode stepping sequence from 1 to 4; C1, C2 represent cur-

rent source and P1, P2 potential measurement. . . . . . . . . . 48

3.6 Impedance mapping system graphical user interface. Displayed

is a typical impedance map. . . . . . . . . . . . . . . . . . . . . 52

3.7 Flow chart of the IMS operational sequence. . . . . . . . . . . . 53

3.8 Measurement made with short circuited electrodes and no mul-

tiplexer front end. . . . . . . . . . . . . . . . . . . . . . . . . . . 54

3.9 Measurement made with short circuited electrodes through mul-

tiplexer front end. . . . . . . . . . . . . . . . . . . . . . . . . . . 55

3.10 RRC test circuit. . . . . . . . . . . . . . . . . . . . . . . . . . . 56

3.11 RRC test circuit attached to the electrode array. All resistor

values (R) are 100 Ω and capacitor value (C) 100 nF. . . . . . . 57

3.12 Typical Cole plots for RRC test circuit, measured data is rep-

resented by red circles and line of best fit by the broken blue

line. (a) Cole plot obtained without the multiplexers (b) Cole

plot obtained with the multiplexers. . . . . . . . . . . . . . . . . 58

3.13 Resistor matrix constructed with 1000 Ω resistors. . . . . . . . . 59

3.14 Typical Cole plot for a measurement made with a resistor ma-

trix, measured data is represented by red circles and line of best

fit by the broken blue line. This Cole plot was obtained from a

matrix constructed of 1000 Ω resistors. . . . . . . . . . . . . . . 61

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3.15 Impedance map obtained from a 1000 Ω resistor matrix, this is

representative of a typical matrix result. (a) R0 map. (b) R∞

map. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 62

3.16 Cole plot of pure plasma. . . . . . . . . . . . . . . . . . . . . . . 63

3.17 Impedance map of pure plasma. (a) R0 map. (b) R∞

map. . . . 64

4.1 Inadequate electrode coverage with bovine blood in the upper

right corner. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 69

4.2 Impedance map with a higher impedance in the upper right

corner due to insufficient electrode coverage. . . . . . . . . . . . 70

4.3 Mean R0 of plasma for various sample thicknesses. . . . . . . . . 72

4.4 Homogeneous haematocrit impedance maps. The haematocrit

is given as a percentage and the colour legend displays the R0

value in ohms. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 74

4.5 Mean R0 for homogeneous impedance maps. . . . . . . . . . . . 75

4.6 Region of red blood cells introduced to plasma. Lower left dark

area is where the cells were injected and the light grey is the

area of visible diffusion. . . . . . . . . . . . . . . . . . . . . . . . 77

4.7 Impedance map of plasma with red blood cells introduced in

lower left corner. . . . . . . . . . . . . . . . . . . . . . . . . . . 78

4.8 Impedance map of plasma with a larger volume of red blood

cells introduced in lower left corner. . . . . . . . . . . . . . . . . 79

4.9 Dispersion of introduced red blood cells into plasma over 2

minute intervals. (a) Measurement made immediately after the

introduction of RBC. (b) Measurement made at 2 minutes. (c)

Measurement made at 4 minutes. (d) Measurement made at 6

minutes. (e) Measurement made at 8 minutes. . . . . . . . . . . 81

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4.10 Impedance map of plasma with introduced red blood cell clot

covering central electrode. . . . . . . . . . . . . . . . . . . . . . 82


covering the 4 electrodes associated with the region in the mid-

dle lower right. . . . . . . . . . . . . . . . . . . . . . . . . . . . 83


covering the 2 lower right regions. . . . . . . . . . . . . . . . . . 84

4.13 Anomalous result due to positive and negative sensitivity fields.

Anomalous result is identified by arrows. . . . . . . . . . . . . . 86

4.14 Reciprocal nature of the tetrapolar electrode configuration. The

reciprocal pairs are grouped . . . . . . . . . . . . . . . . . . . . 88

4.15 Impedance map of plasma used to demonstrate the reciprocal

nature of the tetrapolar electrode configuration. . . . . . . . . . 89

4.16 New proposed electrode stepping sequence 1-8. Red/black and

white/blue represent drive and measurement electrodes respec-

tively. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 91

4.17 Comparison of (a) new electrode stepping sequence with (b)

presently used. . . . . . . . . . . . . . . . . . . . . . . . . . . . 93

4.18 Boundary identification via anomalies. . . . . . . . . . . . . . . 94

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List of Tables

1.1 Cell dimensions in µm used in models of normal tissue (Walker

et al., 2000). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14

1.2 Electrical properties used in cell models (Walker et al., 2000). . 15

2.1 Tissue electrical properties as given by Brown et al, 2000. . . . . 19

4.1 R0 mean and variance for various electrode orientation combi-

nations. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 87

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Statement of original authorship

The work contained in this thesis has not been previously submitted for a de-

gree or diploma at any other higher educational institution. To the best of my

knowledge and belief, the thesis contains no material previously published or

written by another person except where due reference is made.

Jye Geoffrey Smith

xxiii

Acknowledgements

Firstly I would like to thank my supervisor Assoc. Prof. Brian Thomas, or

as he is better know BJ. Without your support and guidance throughout this

project it would not have been possible. You never lost faith even when I

had and our meetings always left me motivated and ready to tackle the next

problem. If I had my time again I promise to use Word instead of LaTeX to

ease what seemed like the endless task of editing.

Assoc. Prof. Bruce Cornish, I am grateful for your help in problem solving,

not only academic but also administrative. Your support and friendly banter

made our meetings a pleasure.

Without the help of Elizabeth Stein and Margaret McBurney I may still

be stuck in a pile of administrative paper work, thankyou.

To everyone I have met, borrowed equipment from, has helped with ex-

periments and even the use of LaTeX throughout this journey I would like to

thank. However insignificant you feel it may have been you have made this

experience a pleasure.

To my darling wife Cheryl, it seemed this day would never come but your

xxv

endless support and encouragement over the years has gotten me to where I

am today.

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List of Abbreviations

BIA Bioimpedance Analysis

MFBIA Multifrequency Bioimpedance Analysis

BIS Bioimpedance Spectroscopy

FEA Finite Element Analysis

CIN Cervical Intraepithelial Neoplasia

AC Alternating Current

DC Direct Current

IMS Impedance Mapping System

RBC Red Blood Cell

xxvii

xxviii

Chapter 1

Introduction

1.1 Cervical Cancer

Cancer of the cervix is one of the most preventable and curable of all cancers

(Greenlee et al., 2000). In 2002 it was the second most common cancer in

women worldwide and it is estimated that up to 90 percent of the most common

type of cervical cancer (squamous cell carcinoma) may be prevented if cell

changes are detected and treated early (Gauthier et al., 1985, Hartikainen

et al., 2001, Koss 1989, Kottmeier 1961, Larsen 1994, Majeed et al., 1994,

Sasieni et al., 1995). Worldwide in 2002 more than 490,000 new cases of

cervical cancer were diagnosed with a mortality of over 270,000 (Parkin et al.,

2007). It is important to note that the death rate from cervical cancer amongst

developing countries is nearly 6 times that of developed countries (Parkin et

al., 2007). There are many reasons for these high death rates but a significant

reduction could be achieved if a device similar to that proposed in this study

1

2 CHAPTER 1. INTRODUCTION

was available for use in screening programs.

Cancer of the cervix affects the cells lining the cervix, which is the lower

part of the womb (uterus) as it joins the inner end of the vagina (Morrow and

Curtin 1998, Nguyen and Averette 1999). Like other cancers, cervical cancer

is a disease where normal cells change, begin to multiply out of control, and

form a growth, tumour or lesion. If not treated early, the growth can invade

local tissue and spread or metastasise to other parts of the body (Boyce et al.,

1984). The main symptoms of cervical cancer are unusual bleeding from the

vagina, and sometimes an unusual vaginal discharge (Raymond et al., 2001).

A cervical cancer may take 10 or more years to develop, but before this

the cells may show pre-cancerous changes. Brown et al. (2000a) and Quek

et al. (1998) proposed that bioimpedance analysis has the potential to detect

these changes and with early treatment there is an excellent chance of a full

recovery. Pre-cancerous lesions can be categorised into two levels of severity,

low-grade abnormalities and high-grade abnormalities, with the higher grade

lesions more likely to progress to a cancer. These are usually graded from warty

atypia (HPV effect), atypia, equivocal cervical intraepithelial neoplasia (CIN),

possible CIN, endocervical dysplasia NOS, CIN1 to CIN3, and carcinoma in

situ. Pre-cancerous changes are relatively easily treated and are cured in nearly

all cases (Baggish (1983)). The type of treatment depends on whether the

change observed is low or high grade, the woman’s age, general health, whether

she wants to have children, and her preferences (Houlard et al., 2002, Raymond

et al., 2001).

3

1.2 Bioelectrical Impedance Analysis

1.2.1 Bioimpedance Background

The Biological Impedance Analysis (BIA) method is an attractive tool as it

offers many opportunities for non-invasive assessment of human body com-

position and tissue characterisation in clinical investigation and patient care

(Baker 1989, Brodie et al., 1998, Chauveau et al., 1999, Rigaud et al., 1994,

Shinkarenko and Kostromina 1998). Advantages such as the non-invasive na-

ture, straightforward use, low cost, safety of operation and high level of re-

producibility provide a rationale for the application of this method (Lukaski

1999).

The use of electricity in medicine began with the discovery that electri-

cal sparks stimulated muscle contraction. This finding prompted investigators

to examine other biological responses to administered electrical current. Al-

though electricity was originally portrayed as a potential therapeutic method,

the measurement of body impedance was subsequently proposed as a use-

ful diagnostic indicator for the physician (King 1970). This hypothesis led

to the identification of physical maladies depicted as low impedance illnesses

and other conditions characterized as high impedance conditions (King 1970),

and has since been advanced to the detection of differing stages of cancerous

growth.

BIA is based upon the relationship between the volume of the conductor

(i.e. the human body or part there of), the conducting length, the compo-

nents of the conductor (e.g., fat or fat free mass, cellular composition) and


its impedance (Z) (Azcue et al. (1993)). Impedance itself reflects frequency-

dependent opposition to the flow of an alternating current, and comprises

resistive (R) and reactive (Xc) components. This impedance is frequency de-

pendent and is defined by equation 1.1 (Ackmann and Seitz 1984).

Z =√

R2 + X2c (1.1)

Both R and Xc components are found in biological systems, although Xc

is usually very small relative to Z at lower frequencies, <4 % (Baumgartner

1996).

It is recognized that the measurement of biological impedance is influenced

by other factors that should either be controlled or reported (Lukaski 1999).

The electrode configuration such as bipolar or tetrapolar is one important

factor (Lukaski 1993, Patterson et al., 1988, Ross et al., 1990).

Electrical conduction in biological systems is mainly ionic, and proportional

to fluid volume and the number of free electrolytic ions (Khaled et al., 1988).

It is also inversely proportional to temperature (Geddes and Baker 1989). This

infers that the bioelectrical resistance is affected by changes in body geometry,

volume, temperature, and electrolytic concentration, and these effects should

be taken into consideration (Geddes and Baker 1989, Khaled et al., 1988).

Biological tissues are inhomogeneous and have quite a complicated struc-

ture containing cells of different shapes and sizes in an extracellular fluid which

behaves as a pure resistance. The cell membranes act as capacitors of about

10 µF cm−2 for muscle cells and about 1 µF cm−2 for other cells, while intra-

cellular fluid presents a resistance to the flow of the electrical current (Kanai

5

1992).

As stated previously, the electrical properties of tissue depend on frequency

and are normally divided into three regions of dispersion: α-dispersion occurs

at low frequencies and is mainly affected by the ionic environment surrounding

the cells; β-dispersion which is a structural relaxation in the frequency range 1

kHz -10 MHz; and at higher frequencies the γ-dispersion which is caused by the

relaxation of the water molecules (Pethig 1987). For many applications the

α- and β-dispersion regions are particularly interesting, since most changes

between normal and pathological tissue seem to appear in these frequency

ranges. Furthermore it is more practical to design a system dedicated to the

low frequency measurement of cancerous tissue (Blad and Baldetorp 1996).

1.2.2 Biological Tissue Equivalent Circuit

The frequency-dependent electrical behaviour of a cellular medium is deter-

mined by the resistivity of the intra-cellular space, the volume of surrounding

extra-cellular fluid, and the capacitance of the membrane (Cole and Cole 1941,

Fricke and Morse 1925, Schwan 1957). The dimensions, internal structure, ar-

rangements of the constituent cells and tissue composition will thus determine

the complex electrical impedance of tissue. This is the basis for the general

application of BIA to assessment of the physiologic ’state’ of organs, and the

cervix (Walker et al., 2000, Walker et al., 2001b).

Significant distinction between normal and precancerous tissue has been

obtained in terms of the low-frequency resistance R0, where R0 is equal to the

extracellular resistance RE , and the high-frequency resistance R∞

, where R∞

is


RE

CRI

Figure 1.1: Schematic of modelled tissue when constructed of its basic com-ponents.

the total resistance at infinite frequency. This may be seen by considering the

electrical analogue RRC circuit for tissue as shown in figure 1.1. Biologically

this is represented in figure 1.2, at low frequencies the current can not pass

through the cell membrane due to it having a high impedance and must flow

through the extracellular space. As the frequency increases the cell membrane

impedance decreases allowing the current to flow through the intracellular

space, hence decreasing the measured tissue impedance.

(a) (b)

Cell Membrane Extracellular Space Intracellular Space

Figure 1.2: (a) Current pathways at low frequencies. (b) Current pathways athigh frequencies.

The impedance spectrum gathered from multifrequency bioimpedance anal-

7

ysis (MFBIA) can be plotted as reactance against resistance. This impedance

spectrum or locus is known as a Cole-Cole plot (figure 1.3), and represents the

impedance of the tissue for frequencies from zero to infinity (Cole and Cole

1941). The impedance at a particular frequency can be represented on the plot

by the phase angle between the vector and the R axis and the magnitude of the

impedance, represented by the length of the vector. Values for RE (extracel-

lular resistance), RI (intracellular resistance) and C (membrane capacitance)

may be deduced from the locus of the Cole-Cole plot and analogous circuit.

R0 as determined from the Cole-Cole plot is equal to the extracellular

resistance RE; and the resistance at infinite frequencies, R∞

, is the parallel

addition of RE and RI as given by equation 1.2.

R∞

=RIRE

RI + RE

(1.2)

IncreasingFrequency

Z

PhaseAngle

R∞

Ro

φ

Impedance atCharacteristic Frequency

Resistance (Ω)

-Reacta

nce

(Ω)

Figure 1.3: Impedance spectrum for modelled tissue when constructed of itsbasic components.

The Cole-Cole plot shown in figure 1.3 is more accurately represented by

equation 1.3 as given by Schwan and Kay (1957). This equation describes the


physical measurements of biological tissue which result in a depression of the

centre of the Cole-Cole semi-circle as shown in figure 1.4, where the centre is

not on the real axis. This depression is a result of varying cell size, structure

and type causing it to be an imperfect capacitor and to have a distribution of

time constants (Cole 1968, Fricke 1932, Schwan 1957).

Zf = R∞

+R0 − R

∞

1 + jωτ 1−α(1.3)

α(π/2)

R∞

Ro

Resistance (Ω)

-Reacta

nce

(Ω)

Figure 1.4: Impedance spectrum for biological tissue with a depressed semi-circle.

At low frequencies, the capacitive cell membranes have high impedance,

and current in cervical tissue is confined to the narrow extracellular pathways

of the highly structured epithelium. This results in a high electrical resistance

(Brown et al., 2000a, Walker et al., 2000). However, in pathological tissue,

these pathways are significantly wider (White and Gohari 1984). In addition,

the reduction in cell volume reduces the most difficult pathways around the

highly flattened cells present in normal epithelium and increases the extracel-

lular space. The expected effect of these changes is a large reduction in low

9

frequency impedance and this suggests that electrical impedance techniques

may provide a method of distinguishing between normal and abnormal cervi-

cal tissue structure (Brown et al., 2000a, Walker et al., 2000).

1.2.3 Applications of BIA to Tissue Characterisation

Cancerous Tissues

Cancer cells in culture show an increased nucleus-to-cytoplasm ratio charac-

terized by increased nuclear size, enlarged nucleoli, and irregular chromatin

distribution (Anderson et al. (1992), Assenheimer et al. (2001) , Blad (1998),

Gonzlez-Correa et al., 1999, Keshtkar et al., 2001, Raymond et al., 2001). In

situations concerning the cervix, cancerous cells appear to be more rounded in

overall shape in contrast to non-cancerous cells, which are flattened along the

growth surface and present a highly structured brick like pattern (Raymond et

al., 2001, Walker et al., 2001a). Figure 1.5 displays the rounded appearance of

cancerous cells that is linked to changes in the structural organisation of actin

polymers (Colgan et al., 2001, Raymond et al., 2001).

Initial clinical trials on an electrical impedance method of diagnosis for

cervical neoplasia have already been undertaken in Sheffield (Brown et al.,

2000a) and by Quek et al. (1998). However Quek also makes use of optical

spectroscope along with bioimpedance. The results of Brown et al. (2000a)

reported clear colposcopy results and good impedance data for 756 measure-

ments made on 124 women. From the data 236 measurements were rejected

due to the tissue not clearly being identified by biopsy or by colposcopy, or

on technical grounds. From comparison of colposcopic and histological results,


Figure 1.5: Changes in tissue structure associated with the progression of CINin cervical squamous epithelium (Walker et al., 2000).

there were 370 measurements from normal squamous epithelium, one from an

invasive cancer, 126 from CIN2/3 (high grade), 63 from CIN1 (low grade),

64 classified as mature metaplasia, 98 classified as immature metaplasia, and

34 classified as columnar tissue. The results demonstrated that BIA provided

significant distinction between normal and precancerous tissue (Brown et al.,

2000a, Mould et al., 1997).

The clinical trials of Brown et al. (2000a) used a tetrapolar configuration

but did not indicate any research conducted towards determining whether

this was the optimum electrode configuration for sampling the required tissue

volumes, even though correct electrode configuration is vital for application

of BIA. The tetrapolar probe utilised by Brown et al. (2000a) provided a

check on small samples of tissue and hence did not provide a impedance map

of an ‘area’ of the cervical tissue. The probe was required to be moved if

additional sampling was required. Quek et al. (1998), although utilising a

different electrode configuration, also only used a single point measurement

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11

system and again the probe was required to be moved to sample additional

areas of the tissue surface.

Wound Mapping

Although developed for a different purpose, a method for monitoring the heal-

ing process of wounds involves a similar concept to that of the current research.

The method was introduced by McCullough et al. (2004) at the XII Interna-

tional Conference On Electrical Bioimpedance & V Electrical Impedance To-

mography. A tripolar electrode configuration implemented into a bandage and

applied to the wound was proposed, with one of the drive electrodes placed onto

the back during measurement. This allowed for the wound to be monitored

over time without the need for the bandage to be removed for examination.

No clinical or experimental data has been presented or published to date.

Breast Cancer

The T-SCAN developed by Assenheimer et al. (2001), and currently approved

by the Food and Drug Administration (FDA), produces two-dimensional maps

of breast tissue impedance via the detection of electrical currents at the surface

of the breast tissue. The T-SCAN injects a current between the patients

breast of interest and arm. Changes in tissue impedance are detected via

the use of an electrode array applied to the breast. These measured changes

are a result of bulk spatial inhomogeneities and may be used to discriminate

between various pathological states. The publication does not state the spatial

resolution required for the early detection of breast lesions.


1.2.4 Tetrapolar Electrode Configuration

The tetrapolar electrode configuration has been accepted as one standard in

virtual biopsy, which is the use of bioimpedance in tissue characterisation.

The tetrapolar surface electrode configuration, consists of two drive electrodes

between which is injected a current, and two measurement electrodes that

record the potential (see Figure 1.6). The ratio of the measured potential to

the amplitude of the imposed current is used to determine the tissue impedance

(Brown et al., 2000).

Drive Electrodes

Measurement Electrodes

Figure 1.6: Tetrapolar arrangement used for taking of bioelectrical impedancemeasurements.

Different electrode configurations have been utilised, for example a tripolar

configuration (McCullough et al., 2004). However the tetrapolar configuration

is that most commonly used. Major advantages of the tetrapolar configuration

are the significant reduction in electrode artefacts (Ragheb et al., 1992), and

that the effect of electrode impedance is removed (Gersing 2001). However

the tetrapolar configuration has a complex pattern of sensitivity within the

13

conductor volume, an important consideration in virtual biopsy since it is

important to know which volume of tissue is being sampled

The relatively few reports of the application of BIA to virtual biopsy have

noted the importance of electrode configuration but have not investigated what

volumes of tissue are being measured. This study will investigate this essen-

tial part of virtual biopsy by means of finite element analysis to model the

tetrapolar configuration.

1.3 Finite Element Analysis (FEA)

Although FEA has a relatively long history in such disciplines as structural

mechanics (Akin 1986, Kardestuncer and Norrie 1987, Holland 1974, Strang

and Fix 1973), its application to field problems in physiology is very recent

(Miller and Henriquez 1990, Pavlin et al., 2001, Radai et al., 1999). The 1982

review article by Heringa et al. (1982) discussed only three methods of solution

for bioelectrical field problems, one of which was the finite element method.

One of the first papers reporting application of FEA to modelling of electrical

characteristics of tissue with reference to virtual biopsy was by Kanai (1992),

with work on frequency characteristics of electrical properties of living tissue

and its clinical applications.

Walker et al. (2000) have used the FEA technique to construct models of

the different cells that occupy multiple layers in normal or cancerous epithelium

of the cervix. A cross-section of a modelled single cell used by Walker et al.

(2000) is shown in Figure 1.7. The cells dimensions were assigned from the


values given by (Friedrich and Morse, 1925), and electrical properties based

on conductivities and capacitance quoted by Irimajiri et al., (1978) and Pethig

(1984). These properties are shown in Tables 1.1 and 1.2 respectively.

CytoplasmExtracellular

Space

Nucleus

(a) (b)

40mm

40mm

Electrodes

Stroma (5mm)

Basement

Membrane (100nm)

Epithelial

Layer (0.3mm)

Figure 1.7: (a) Cross section of finite element model of single cell. (b) Finiteelement tissue model (Walker et al., 2000, 2001a, 2001b)

In the studies of Walker et al. the models suggested that current flow

was confined to the epithelium. However these results did not identify which

volumes of tissue produced the majority of the impedance spectrum detected.

There is an apparent need to investigate these volumes such that the tissues

sampled are those of importance to the biopsy required.

Table 1.1: Cell dimensions in µm used in models of normal tissue (Walker etal., 2000).

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15

Table 1.2: Electrical properties used in cell models (Walker et al., 2000).

1.4 Aims and Objectives

The overall objective of this research was to design and construct an impedance

mapping system (IMS) for objective virtual biopsy of lesions by bioimpedance

spectroscopy (BIS). Initially freshly excised cervical tissue was to be tested

however as the study progressed this proved problematic and bovine blood

was chosen as a suitable substitute.

Specific aims were to; Investigate the spatial sensitivity distribution of the tetrapolar electrode

configuration via finite element analysis (FEA). Design a novel front end multiplexing system and multi-electrode array

for mapping the impedance of the tissue of interest. Experimentally confirm the efficacy of the approach to identify regions of

different impedances and their boundaries using bioimpedance mapping.

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Chapter 2

Modelling of Sensitivity

Distributions

2.1 Introduction

The tetrapolar electrode configuration is commonly used for tissue characteri-

sation by bioimpedance and involves injecting a constant drive current between

an adjacent pair of electrodes (drive electrodes), and measurement of the re-

sulting potential between another pair of adjacent electrodes (measurement

electrodes). This measured potential is dependent on the electrical charac-

teristics of the volume of tissue under investigation. This study looks at how

various volumes of tissue and their location with respect to the electrodes affect

the resulting measurement.

Finite element analysis (FEA) was used to model the sensitivity of the

tetrapolar electrode configuration with an emphasis on the sensitivity for de-

17

18 CHAPTER 2. MODELLING OF SENSITIVITY DISTRIBUTIONS

tecting a cancerous lesion on the surface of healthy cervical tissue. The elec-

trodes were modelled in contact with the surface of the tissue. It was found that

the sensitivity varied with differing lesion locations, and when modelled with

the lesion located between a drive and measurement electrode pair, anomalous

results were obtained. The existence of the anomaly was confirmed experi-

mentally.

Sensitivity Fields for the tetrapolar configuration were also modelled in 3

dimensions using the approach as described by Grimnes and Martinsen (2006).

The model confirmed the existence of the anomaly and provided additional in-

sight into current density and sensitivity as a function of depth for the tetrap-

olar configuration.

2.2 Lesion Modelling

2.2.1 Methodology

FEA

The FEA method had been previously validated during research undertaken as

part of the Bachelor of Applied Science Honours degree. The tetrapolar elec-

trode configuration was modelled on physical dimensions that are achievable

when constructing probes (Brown et al, 2000). Electrodes were modelled as 1

mm in diameter and separated by 0.2 mm as shown in figure 2.1. Electrodes

were given ideal electrical properties of zero resistivity and contact impedance

with the modelled tissue. As the interest of this study was the volume of tis-

19

sue on and just below the surface, the lesion was modelled as a hemisphere

located immediately below the surface of healthy tissue. Two arrangements

were modelled; (a) with a lesion central to the electrode configuration and

(b) with the lesion situated between a drive and measurement electrode. This

latter arrangement is shown in figure 2.1.

Tissues were assigned mean electrical properties as described in Brown et

al, 2000 (Table 2.1) and applied by equation 2.1 (Schwan and Kay 1957).

Where R0 = RE and R∞

and Fc were calculated by equations 1.2 and 2.2

respectively. In this equation Fc is the characteristic frequency and defined

as the frequency where the reactance is greatest. The frequency of the drive

current is F and for the purpose of this study α (figure 1.4) was given the value

of zero since the cell membrane was modelled as a perfect capacitor.

Z = R∞

+(R0 − R

∞)

(1 + [jF/Fc])1−α(2.1)

Fc =1

2πC(RE + RI)(2.2)

Table 2.1: Tissue electrical properties as given by Brown et al, 2000.

The lesion was modelled with various radii from 0 mm (no lesion present,

only healthy tissue) to 1.05 mm. For the lesion modelled central to the elec-

trodes a direct current (DC) was injected and the potential between mea-

surement electrodes calculated. For the lesion located between a drive and

halla



Drive

Electrodes

Measurement

Electrodes

Lesion

mm

mm

(a)

(b)

Figure 2.1: (a) Tetrapolar electrode model with the lesion located betweena drive and measurement electrode. (b) 3D view showing the lesion locatedimmediately below the surface of healthy tissue.

21

measurement electrode, the modelling was expanded to comprise a frequency

spectrum of 600, 1200, 2400, 4800, 9600, 19200, 38400, 76800, 153600, 307200

and 614400 Hz and the real part of the potential between measurement elec-

trodes determined. The magnitude of the drive current was kept constant

throughout.

Experimental Verification

Experimental verification of the modelled tetrapolar electrode configuration

was difficult due to the small size of the electrodes and lesion. For this reason

the electrode size was altered but the same tetrapolar configuration main-

tained. The experimental verification was performed for the case of a lesion

located (a) central to the electrodes and (b) between a measurement and drive

electrode i.e. where the modelling indicated an anomalous result for the mea-

sured impedance. The experimental setup is shown in figure 2.2.

50 mm

50

mm rDrive Electrodes

(Luer needles)

Measurement Electrodes

(Luer needles)

(a) (b)

Figure 2.2: Experimental electrode placement on the surface of a saline solutionand brass rod with radius ’r’ located (a) central to the electrodes (b) betweena drive and measurement electrode.

The electrodes (16 G (1.2 mm) Luer needles) were separated by 50 mm (in a

50 mm square) and held in place with a non-conducting epoxy. The electrode


configuration was then placed on the surface of a saline solution with the

needles penetrating 4 mm into the solution. The saline solution represented

a large medium of healthy tissue. Brass rods of various diameters were used

to represent lesions of lower resistance compared to healthy tissue (saline).

The rods were 30 mm in length and varied in radius from 2.5 mm to 22.5

mm. They were placed upright in the saline solution just below the surface. A

SEAC SFB3 body composition meter was used to supply a constant current

at frequencies of 10, 15, 20, 25, 35, 55, 65, 80, 120 and 310 kHz between the

drive electrodes and the potential between measurement electrodes recorded

when each rod was positioned as indicated above.

The experimental setup was also modelled using the FEA method. The

model included dimensions as described above in the experimental setup and

the conductive mediums were given electrical properties of saline and brass.

2.2.2 Results

FEA

All potentials were normalized against the potential measured for homogeneous

healthy tissue at DC (0 Hz). For a lesion located central to the electrodes the

potential decreased as the radius of the lesion with lower impedance increased

(figure 2.3). However for a lesion located between a drive and measurement

electrode the results (figures 2.4 & 2.5) clearly demonstrate a rise in measured

potential (particularly at the lower frequencies) as the lesion increased in radius

up to approximately 0.4 mm followed by a decrease in potential. Shown in

figure 2.6 is the potential versus frequency for (a) healthy tissue and (b) healthy

23

tissue with a lesion 0.75 mm in radius located between a drive and measurement

electrode.

The detection of precancerous changes using BIA relies upon detecting the

reduced impedance of these changes at low frequencies compared with that

of healthy tissue. The expected effect was noticed when a region of lower

impedance was introduced into the centre of the tetrapolar electrode configu-

ration. The results of the FEA modelling displayed in figure 2.3 show that a

lesion of 0.4 mm radius reduced the measured impedance by 30 % and a 0.8

mm lesion would approach an impedance measured for a lesion alone. How-

ever an increase in measured potential was noted in the modelled results with

the introduction of a region of lower impedance midway between a drive and

measurement electrode. In the modelled results (figure 2.4) the increase at DC

was a maximum for a lesion of radius 0.4 mm. For larger radii the impedance

decreased as expected but did not reach the impedance for a lesion alone until

a radius greater than 1.0 mm. This anomaly was absent in the case of a lower

impedance medium central to the electrodes. This anomaly could cause a pre-

cancerous lesion to be identified as healthy tissue as demonstrated in figure 2.6

where results with a relatively large lesion of 0.75 mm in radius situated be-

tween a drive and measurement electrode are similar to that of healthy tissue.

Experimental Results

All potentials for the experimental and modelled experimental results were

normalized against their respective potentials measured for the pure saline


0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.80.4

0.5

0.6

0.7

0.8

0.9

1

1.1

Lesion Radius (mm)

Nor

mal

ised

Pot

entia

l

Figure 2.3: Modelled results of tetrapolar electrode configuration for a lesionlocated centrally in the electrode configuration.

solution. Results for a centrally located brass rod (representing a lesion) are

shown in figure 2.7. The measured potential is seen to decrease as the radius of

the rod increased in a similar manner to that indicated by the FEA modelling.

Again as predicted by the FEA modelling, the experimental results show an

initial rise in measured potential as the radius of the brass rod increased when

the rod was located between a drive and measurement electrode (figure 2.8).

Figure 2.8 also shows the results of FEA modelling of the experimental

electrode configuration. The position of the maximum potential is similar in

the experimental and modelled results. However the magnitude of the potential

is consistently less than the experimental results for both brass rod locations

(central to the electrodes and between a drive and measurement electrode).

The agreement between the modelled and experimental results in indicating

the position of the maximum provides confidence that the results of FEA

25

0 0.2 0.4 0.6 0.8 1 1.2 1.40

0.5

1

1.5

Lesion Radius (mm)

Nor

mal

ised

Pot

entia

l

00.61.22.44.89.619.238.476.8153.6307.2614.4

Figure 2.4: Modelled results of tetrapolar electrode configuration for a lesionlocated midway between a drive and measurement electrode (Legend displaysfrequency in kHz).

10−1

100

101

102

103

0

0.5

1

1.5

Frequency (kHz)

Nor

mal

ised

Pot

entia

l

0.000.050.150.250.350.450.550.650.750.850.951.05

Figure 2.5: Modelled results of tetrapolar electrode configuration for a lesionlocated midway between a drive and measurement electrode (Legend displayslesion radius in mm).


10−1

100

101

102

103

0.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

1.1

Frequency (kHz)

Nor

mal

ised

Pot

entia

l0.00 mm0.75 mm

Figure 2.6: Measurement with (a) no lesion and (b) a lesion of radius 0.75 mmin otherwise healthy tissue.

modelling are realistic.

0 5 10 15 200.1

0.2

0.3

0.4

0.5

0.6

0.7

0.8

0.9

1

1.1

Lesion Radius, mm

Nor

mal

ised

Pot

entia

l

ExperimentalModelled

Figure 2.7: Experimental and modelled results of the tetrapolar electrode con-figuration for a brass rod located centrally in the electrode configuration.

27

0 5 10 15 200.9

1

1.1

1.2

1.3

1.4

1.5

Lesion Radius, mm

Nor

mal

ised

Pot

entia

l

ExperimentalModelled

Figure 2.8: Experimental and modelled results of the tetrapolar electrode con-figuration for a brass rod located midway between a drive and measurementelectrode.

2.3 Sensitivity Fields

Sensitivity fields have been used to describe the tetrapolar configuration pre-

viously by Brown et al., 2000 and more recently by Grimnes and Martinsen,

2006. However in neither case were the electrodes modelled in a realistic square

configuration for tissue characterisation applications; Brown et al. modelled a

square tetrapolar electrode configuration using point electrodes and Grimnes

and Martinsen modelled using a line configuration. In the present study a real-

istic square electrode configuration was modelled. The geometry and electrode

properties were the same as previously described (see section 2.2.1) with the

exception that the electrodes had a diameter of 1 mm. The electrode spacing

was varied and the sensitivity determined as a function of depth into the test

medium.


The sensitivity of a volume of medium is a measure of how much this volume

contributes to the total measured impedance. Geselowitz (1971) showed that

the sensitivity, S, at a point within the model can be defined as the scalar

product of the vector current densities (equation 2.3). Here the current density

J1 at any point is found by injecting the current, I, between the drive electrodes.

J2 is the current density found by now injecting the same current between the

measurement electrodes.

S =J1 · J2

I2(2.3)

The resultant field can have volumes of positive and negative sensitivity. If

positive then a drop in measured impedance will result if a medium of lower

resistivity is located in this volume. Whereas a negative value would result

in an increase in measured impedance. The measured impedance will likewise

increase and decrease for positive and negative sensitivity values respectively

when a medium with a higher resistivity is introduced. The change in measured

impedance is also dependent on the magnitude of the sensitivity value.

2.3.1 3-D Sensitivity Field Model

Modelling used the same electrode dimensions as previously described (1 mm

diameter electrodes) with a larger conductive medium to reduce the effect of

confined current flow on the results. A confined current flow would distort the

current density as demonstrated in figure 2.9 resulting with incorrect sensitivity

fields. Sensitivity fields were also found for models with an electrode spacing of

0.4 mm, 0.6 mm, 0.8 mm and 1.0 mm to investigate the relationship between

29

electrode spacing and measurement depth. The injected current is between

the upper and lower left electrodes and has an amplitude of 1 A.

Figure 2.10 through to figure 2.15 display the sensitivity fields for a model

with an electrode spacing of 0.2 mm and for a depth of 0.0 (surface of the

conductive medium), 0.2, 0.4, 0.6, 0.8 and 1.0 mm. The electrode positions on

the surface are represented in the figures by black circles.

The fields show very localised areas of positive sensitivity between the drive

or measurement electrode pairs and more importantly negative sensitivity be-

tween drive and measurement electrode pairs. These areas of negative sen-

sitivity become significantly larger with increasing depth but the magnitude

decreases substantially (a factor greater than 10 at 0.2 mm depth). The area

of positive sensitivity (shown in red on figures 2.10 - 2.15) also increases in

size with depth but not to the same extent. The lack of sensitivity on the sur-

face central to the electrodes explains the relatively insensitive nature of the

measurement for small lesions (see figures 2.3 and 2.7). The lesion must be of

sufficient size to overlap regions of sensitivity deeper into the tissue and to the

sides before being detected. Negative fields between the drive and measure-

ment electrode pairs agree with the results obtained from modelling lesions and

experimental verification. This negative sensitivity field shows that a lesion of

lower resistance located here will result in an increase in measured impedance

(i.e. anomalous results).

The maximum of the sensitivity over the plane at each depth is plotted in

figure 2.16 and the sum of the absolute sensitivity in the plane in figure 2.17.

It can be seen that the maximum sensitivity decreases significantly at a depth

of 0.1 mm and is less than 10 % of the maximum at 0.2 mm. The sum of the


D

D

Figure 2.9: 3-D modelled tetrapolar configuration and current density confinedby the boundaries.

5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

-0.4

-0.3

-0.2

-0.1

0

0.1

0.2

Figure 2.10: Sensitivity field at a depth of 0.0 mm.

31

5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

-0.02

-0.015

-0.01

-0.005

0

0.005

0.01


5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

-6

-4

-2

0

2

4

6

x 10 -3



5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

-2.5

2-

-1.5

-1

-0.5

0

0.5

1

1.5

2

2.5

x 10 -3


5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

-1

-0.5

0

0.5

1

x 10 -3


33

5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

-4

-2

0

2

4

6

x 10 -4


absolute sensitivity over the planes with respect to depth approaches a value

less than 10 % of the original for depths greater than 0.4 mm.

Sensitivity fields have confirmed the existence of the anomaly (negative

sensitivity) within the tetrapolar electrode configuration and that it is confined

to the regions between drive and measurement electrode pairs. It can also be

seen that the sensitivity fields indicate that the measured impedance is limited

to the surface layers, predominantly to a depth of less than 0.4 mm, making

the tetrapolar configuration ideal for detection of surface lesion.

Grimnes and Martinsen (2006) note that equation 2.3 demonstrates the

reciprocal nature of the tetrapolar configuration. Under linear conditions (e.g.

same drive and measurement electrode size) the drive and measurement elec-

trode are interchangeable without a change in measured impedance. This is

also shown in the symmetrical nature of the sensitivity fields since the same


0 0.2 0.4 0.6 0.8 10

0.05

0.1

0.15

0.2

0.25

0.3

0.35

Depth, mm

Sen

sitiv

ity

0.2 mm0.4 mm0.6 mm0.8 mm1.0 mm

Figure 2.16: Maximum sensitivity in each plane as a function of depth forelectrode spacing’s 0.2 mm, 0.4 mm, 0.6 mm, 0.8 mm and 1.0 mm.

sensitivity field is produced independent of which electrode pair (left or right)

are the drive electrodes.

Later in chapter 4 it will be shown that the anomaly produced by negative

sensitivity is present in the impedance maps and that it may be noted for

further analysis.

Electrode Spacing and Measurement Depth

To confirm the use of sensitivity as an indicator of measurement depth, the

tetrapolar configuration was again modelled with a range of electrode spacing’s

and the measurement medium thickness varied. Each electrode spacing was

modelled with a medium thickness of 0.2, 0.4, 0.6, 0.8 and 1.0 mm. The

35

0 0.2 0.4 0.6 0.8 10

2

4

6

8

10

12

14

Depth, mm

Sen

sitiv

ity

0.2 mm0.4 mm0.6 mm0.8 mm1.0 mm

Figure 2.17: Sum of the absolute sensitivity in each plane as a function ofdepth for electrode spacing’s 0.2 mm, 0.4 mm, 0.6 mm, 0.8 mm and 1.0 mm.

resulting potentials were normalised to one and are shown in figure 2.18.

It can be seen that with increasing medium thickness the measured po-

tential decreases until a medium thickness of 5 to 6 mm. This thickness is

significantly different to that given by the maximum or sum of the absolute

sensitivity (0.2 & 0.4 mm). It is later experimentally verified that the thick-

ness must be at least 4 to 5 mm thick so as not to distort the impedance

measurements (see section 4.1.2). This result suggests that sensitivity is not

a good indicator of measurement depth and should only be used for mapping

positive and negative sensitivity fields.


2 3 4 5 6 7 8 9 100

0.2

0.4

0.6

0.8

1

Tissue Thickness, mm

Nor

mal

ised

Pot

entia

l

0.2 mm0.4 mm0.6 mm0.8 mm1.0 mm

Figure 2.18: Measured potential against tissue medium thickness for differingelectrode spacings.

2.3.2 Electrode Array Sensitivity Fields

The 3 dimensional model was expanded to include surrounding inactive elec-

trodes as would be the case for the impedance mapping system (IMS) described

in chapter 4. The electrode spacing was also increased to 0.77 mm which is

representative of the spacing used in the instrumentation later developed. The

maximum and sum of the sensitivity in each plane for a range of depths was

found for each of the four unique electrode sites shown in figure 2.19. All of

the other electrode sites located on the array can be represented by one of the

four by rotation, reflection and taking into account the reciprocal nature of

the electrode configuration.

Figures 2.20 & 2.21 show that the maximum and sum of the absolute sen-

sitivity decrease to 10 % of the original value much deeper into the model (0.6

37

D

D

M

M

D

D

M

M

D D

MM

D

D

M

M

(1) (2)

(3) (4)

Figure 2.19: The four unique electrode sites modelled.

& 0.4 mm respectively) compared to the single tetrapolar electrode model (see

figures 2.16 & 2.17 for comparison). The maximum sensitivity has also de-

creased to 5 % of the original value modelled without the inactive electrodes.

This large decrease indicates that the inactive array is altering the current dis-

tribution, however the sensitivity fields produced (figure 2.22) still show strong

areas of positive and negative sensitivity directly below the active electrodes.

Each of the four unique electrode sites display similar plots for maximum

and sensitivity sum (figures 2.20 & 2.21). Indicating that impedance measure-

ments taken on a homogeneous medium will be similar whether taken with an

electrode set inside or on the border of the array.


0 0.2 0.4 0.6 0.8 10

0.002

0.004

0.006

0.008

0.01

0.012

Depth, mm

Sen

sitiv

ity

Location 1Location 2Location 3Location 4

Figure 2.20: Maximum sensitivity in each plane as a function of depth forelectrode sites 1, 2, 3 and 4 (see figure 2.19).

Electrode Array Measurement Depth

The effect of inactive electrodes on measurement depth was modelled using site

1 (figure 2.19) and an electrode spacing of 0.77 mm which is the experimental

electrode spacing. The medium was increased in thickness from 0.2 mm to

1.0 mm and the results are shown in figure 2.23. The results for a single

tetrapolar electrode configuration of electrode spacing 0.77 mm are also shown

in figure 2.23.

With a small medium thickness the results for the single electrode config-

uration and the array are identical, with the array plot overlaying the plot

for a single electrode. This result shows that the array of inactive electrodes

has little effect on the resultant measurement depth as previously indicated by

the plots of maximum sensitivity and sum of absolute sensitivity (figures 2.16,

39

0 0.2 0.4 0.6 0.8 10

0.2

0.4

0.6

0.8

1

1.2

1.4

Depth, mm

Sen

sitiv

ity

Location 1Location 2Location 3Location 4

Figure 2.21: Sum of the absolute sensitivity in each plane as a function ofdepth for electrode sites 1, 2, 3 and 4(see figure 2.19). Locations 1 and 3 aredifficult to see due to overlap.

5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

14

12

10

8

6

4

2

0

2

D

D M

M

x 10-3

Figure 2.22: Sensitivity field at a depth of 0.1 mm for electrode site 1 shownin figure 2.19.


0 2 4 6 8 100.6

0.65

0.7

0.75

0.8

0.85

0.9

0.95

1

1.05

1.1

Thickness, mm

Nor

mal

ised

Pot

entia

l

Single Electrode ConfigurationArray

Figure 2.23: Measurement thickness for an array with 0.77 mm electrode spac-ing and a single electrode configuration with the same spacing.

2.17, 2.20 & 2.21). Figure 2.23 also shows that the measurement depth is much

greater than that indicated by the plot of sensitivity against depth. Once again

confirming that sensitivity is not appropriate as an indicator of measurement

depth.

Effect Of Inactive Electrodes Inside The Tetrapolar Configuration

The tetrapolar electrode configuration was modelled as an array with inactive

electrodes within the active set to determine if the inactive electrodes would

affect measurements of a larger surface area. Figure 2.24 displays the resultant

field at a depth of 0.1 mm.

It can be seen that the inactive electrodes between the drive and measure-

41

5 10 15 20 25 30 35 40 45 50 55 60

5

10

15

20

25

30

35

40

45

50

55

60

6

4

2

0

2

4

6

8

10

x 10 -4

D

D M

M

Figure 2.24: Sensitivity field at a depth of 0.1 mm for a tetrapolar electrodeconfiguration with inactive electrode inside the configuration.

ment electrodes result in an area of larger negative sensitivity underneath their

location (this may not be clear in the hard copy, please refer to the electronic

copy provided on the CD). Any measurements made with inactive electrodes

within the active set would therefore be affected and could produce misleading

results similar to the anomaly previously found. This is due to the inactive

electrodes being located in regions of negative sensitivity.

However large surface area measurements such as this are not necessary for

the proposed impedance mapping system. If high imaging resolution of the

surface tissue is to be achieved then the electrode spacing must be as small as


possible. Allowing smaller surface areas to be analysed and lesion boundaries

more accurately identified.

Chapter 3

Impedance Mapping System

Overview, Operation and

Testing

Tissue characterisation by bioimpedance analysis has to date been single point,

where a large tissue sample was analysed at a single point and no attempt was

made to identify the boundaries of any lesion. Sampling of a single point also

means that precancerous changes might not be detected if the point chosen for

sampling was healthy tissue and other areas were indeed precancerous. This

study aims to reduce the possibility of not detecting areas of precancerous

change by identifying boundaries of any change detected by means of develop-

ing a novel impedance mapping system (IMS) to sample a large area of tissue.

The IMS will make use of the tetrapolar electrode configuration in a repeated

array allowing measurements to be taken over the surface of a tissue sample.

BIA of individual measurement locations can then be visually displayed and

43

44CHAPTER 3. IMPEDANCE MAPPING SYSTEM OVERVIEW,

OPERATION AND TESTING

boundaries of a suspected lesion identified. This approach of display will also

allow for easy identification of anomalous results which, as discussed in chapter

2, can result if a lesion is located between a drive and measurement electrode.

The IMS developed for this study is shown schematically in figure 3.1. The

current source and potential measurements are switched through multiplexers

which allow them to be directed to any electrode on the array. A PC based

program was written in Visual Basic (VB) to switch the multiplexers and

complete all data analysis through Matlab for the visual display. An array

consisting of 25 electrodes allows 16 measurement sites over the surface of a

medium and 4 tetrapolar measurements at each site. This provides a total of

64 measurements on the area of medium covered by the array.

PC

SFB7 Multiplexers Electrode Array

Figure 3.1: Schematic of the bioimpedance mapping system.

45

3.1 Current Source & Potential Measurement

An Imp SFB7 from Impedimed (www.impedimed.com) was used as the current

source and for potential measurement. The SFB7 is a single channel, tetra po-

lar bioimpedance spectroscopy (BIS) device with 265 frequencies in the range

of 4 kHz to 1000 kHz. The SFB7 has an impedance accuracy of +/- 1.0% in

the range 50 to 1100 Ω and allows access to the full raw data of resistance and

reactance.

3.2 Multiplexers

Current and potential probes from the SFB7 were directed to the appropriate

electrode in the array via 32-channel ADG732 (figure 3.2) multiplexers from

Analog Devices. The ADG732 was chosen as they have a low ‘on’ resistance

(Ron) of 4 Ω through each channel. The Ron resistance can be considered

negligible since when in series with the source probes a constant current is

always driven. Ron is also negligible with the potential probes since they draw

no current and this results in no change to the measured potential.

Multiplexer switching is by 5-bit binary parallel inputs and switching may

be enabled or disabled (switching locked) with the addition of another input.

The parallel inputs and disable function make them ideal for control with a

parallel port found on any desktop personal computer. Figure 3.3 shows the

multiplexer circuit diagram with parallel port (Header 9) and electrode array

socket (Header 25). Pins Red and Black are SFB7 current source connections

and pins White and Blue are SFB7 potential measurement connections.



Figure 3.2: Functional block diagram and pin configuration of multiplexerADG732.

3.3 Electrode Array

The electrode array consisted of 25, 1 mm diameter electrodes separated by

0.77 mm in a 5x5 square as shown in figure 3.4. The use of 25 electrodes instead

of 32, which is the total available multiplexer channels, allowed for the array

to be constructed on a single sided printed circuit board (PCB). The use of a

larger array would require a double sided PCB and further difficulties coupling

with the multiplexers. FEA modelling showed that the measurable depth was

independent of electrode spacing, so the electrodes were positioned as close as

possible to increase the imaging resolution. A PCB edge connector (Header

25, figure 3.3) allows easy removal of the electrodes from the multiplexers.

Using a PCB made manufacturing the electrode arrays cheap and disposable

for single use or if needed they could be autoclaved and sterilised for repeated

uses, which was the original intent.

Measurements were made using the tetrapolar configuration with 4 mea-

47

1

1

2

2

3

3

4

4

5

5

6

6

7

7

8

8

D D

C C

B B

A A

Title

Number RevisionSize

A3

Date: 7/07/2007 Sheet of

File: F:\Latex\Thesis exers4x.SCHDOC Drawn By:

+5 +5

+5

+5

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25

Header 25

pin3

pin4

pin5

pin6

pin7

pin8

pin9

pin1

0

pin1

1

pin1

2

pin1

3

pin1

4

pin1

5

pin1

6

pin1

7

pin18

pin1

pin2

pin3

pin4

pin5

pin6

pin7

pin8

pin9

pin10

pin11

pin12

pin13

pin14

pin15

pin16

pin17

pin18

pin19

pin20

pin21

pin22

pin23

pin24

pin25 pin1

pin2

pin3

pin4

pin5

pin6

pin7

pin8

pin9

pin1

0

pin1

1

pin1

2

pin1

3

pin1

4

pin1

5

pin1

6

pin17

pin18

pin19

pin20

pin21

pin22

pin23

pin24

pin25

pin1

pin2

pin3

pin4

pin5

pin6

pin7

pin8

pin9

pin1

0

pin1

1

pin1

2

pin1

3

pin1

4

pin1

5

pin1

6

pin1

7

pin1

8

pin1

9

pin2

0

pin2

1

pin2

2

pin2

3

pin2

4

pin2

5

pin1

pin2

pin19

pin20

pin21

pin22

pin23

pin24

pin25

pin1

pin2

pin3

pin4

pin5

pin6

pin7

pin8

pin9

pin1

0

pin1

1

pin1

2

pin1

3

pin1

4

pin1

5

pin1

6

pin1

7

pin18

pin19

pin20

pin21

pin22

pin23

pin24

pin25

11

22

33

44

55

66

77

88

99

1010

1111

1212

1313

1414

IN1

15

IN2

16

IN3

17

IN4

18

IN5

19

IN6

20

WRb

ar21

ENba

r22

GND

23

Vss

24

NC25

NC26

NC27

NC28

NC29

NC30

NC31

3232

3333

3434

3535

3636

3737

3838

3939

4040

NC41

NC42

D43

NC44

4545

4646

4747

4848

Multiplexer - ADG732

Top_Driver

Multiplexer

11

22

33

44

55

66

77

88

99

1010

1111

1212

1313

1414

IN1

15

IN2

16

IN3

17

IN4

18

IN5

19

IN6

20

WRb

ar21

ENba

r22

GND

23

Vss

24

NC25

NC26

NC27

NC28

NC29

NC30

NC31

3232

3333

3434

3535

3636

3737

3838

3939

4040

NC41

NC42

D43

NC44

4545

4646

4747

4848


Under_Driver

Multiplexer

11

22

33

44

55

66

77

88

99

1010

1111

1212

1313

1414

IN1

15

IN2

16

IN3

17

IN4

18

IN5

19

IN6

20

WRb

ar21

ENba

r22

GND

23

Vss

24

NC25

NC26

NC27

NC28

NC29

NC30

NC31

3232

3333

3434

3535

3636

3737

3838

3939

4040

NC41

NC42

D43

NC44

4545

4646

4747

4848


Top_Potential

Multiplexer

11

22

33

44

55

66

77

88

99

1010

1111

1212

1313

1414

IN1

15

IN2

16

IN3

17

IN4

18

IN5

19

IN6

20

WRb

ar21

ENba

r22

GND

23

Vss

24

NC25

NC26

NC27

NC28

NC29

NC30

NC31

3232

3333

3434

3535

3636

3737

3838

3939

4040

NC41

NC42

D43

NC44

4545

4646

4747

4848


Under_Potential

Multiplexer

Blac

k

Red W

hite

Blue

dn3

dn4

dn5

dn6

dn7

dn8

dn9

dn3

dn4

dn5

dn6

dn7

cn1

cn2

cn1

cn2

0.1uF

C190.1uF

C18

0.1uF

C17

0.1uF

C16

1 2 3 4 5 6 7 8 9

Header 9

dn3

dn4

dn5

dn6

dn7

dn8

dn9

dn3

dn4

dn5

dn6

dn7

dn3

dn4

dn5

dn6

dn7

Figu

re3.3:

Multip

lexer

circuit

diagram

.Parallel

port

represen

tedby

Head

er9

and

electrode

arrayso

cketby

Head

er25.

SFB

7drive

and

measu

remen

telectro

des

connect

throu

ghred

,black

and

white,

blu

eresp

ectively.



Electrodes

Figure 3.4: PCB electrode array.

P2 C2

P1 C1

P1 P2

C1 C2

C1 P1

C2 P2

C2 C1

P2 P1

(1) (2)

(3) (4)

Figure 3.5: Electrode stepping sequence from 1 to 4; C1, C2 represent currentsource and P1, P2 potential measurement.

49

surements being made for each measurement site on the surface of the test

medium. As shown in figure 3.5 measurements were made by switching the

current and potential electrodes such that the tetrapolar configuration was

effectively rotated by 90o for each successive measurement. The tetrapolar

configuration was then shifted to the right by one column and another 4 mea-

surements taken. This was repeated for all columns and moved down one row

at a time and repeated. This gives a total of 64 separate measurements to

be taken at 16 different locations and an impedance map of 49 mm2 on the

surface of a test medium.

3.4 Impedance Data Analysis

Measurements were made via stepping through the electrode array sequence

as detailed previously and values for resistance and reactance were taken with

the SFB7 and imported to Matlab. A circular locus of best fit to reactance

versus resistance was determine by applying equation 3.1.

R2 + X2− aR − bX + c = 0 (3.1)

R2+X2 can be substituted for Z2 and equation 3.2 solved by multiple re-

gression (Cornish, 1994).

Z2 = aR + bX − c (3.2)

Values a, b and c describe the circle of best fit (equation 3.3 and equa-

tion 3.4) and were used to determine the cole parameters R0 and R∞

by equa-



tion 3.5 and equation 3.6.

centre = (a

2,b

2) (3.3)

radius =

√

a2

2+

b2

2− c (3.4)

R0 =a

2−

√

radius2− (

b

2)2 (3.5)

R∞

=a

2+

√

radius2− (

b

2)2 (3.6)

3.5 Impedance Mapping System Graphical Use

Interface and Operation

The IMS graphical user interface (GUI) shown in figure 3.6 was developed

as part of this study in Visual Basic to control the multiplexer front end

stepping sequence. The GUI was also used as a visual display for the impedance

measurements and mapping of the 16 regions on the surface of the medium

measured. ‘Gather’ starts the mapping process and multiplexers switching.

The array of 25 squares represent the electrode array and change colours to

indicate where the electrode stepping sequence is currently present. Red and

black represent drive electrodes and white and blue measurement electrodes.

When this sequence is finished the ’Data Analysis’ button starts the Matlab

analysis of the raw resistance and reactance data. The 64 R0 values calculated

in Matlab are imported into the GUI and displayed in 4 smaller impedance

51

maps, representing each of the tetrapolar electrode orientations as shown in

figure 3.5. The maps electrode orientation is indicated by red, white, blue

and black squares above each map. The four electrode orientation maps are

averaged to present the larger impedance map. Figure 3.7 shows a flow chart

that more simply presents the IMS GUI operational sequence. The attached

CD contains a copy of the IMS program with a set of real data which is later

displayed in figure 4.7.

Impedance maps may also be displayed for R∞

and capacitance of the test

medium. This feature can be expanded to present any of the Cole or user

defined parameters.

The colour bar legend is scaled for each impedance map so that the range

is always from the smallest to largest R0 (or user defined parameter) value.

Scaling in this manner will insure that the greatest colour contrast possible is

presented in the map.

3.6 Impedance Mapping System Testing

3.6.1 Testing of Multiplexer Ron Contribution

The 32 channel ADG732 multiplexers were chosen due to their low resistance

(Ron) when a channel was in use. Theoretically this resistance should play

no part in measurements due to the resistance being in series with the drive

electrodes which pass a constant current and in series with the measurement

electrodes which draw no current. To test this hypothesis short circuited mea-



Figure 3.6: Impedance mapping system graphical user interface. Displayed isa typical impedance map.

surements (no impedance load used) were made in the frequency range 4 kHz to

1000 kHz with and without the multiplexer front end. Measurements without

and with the multiplexers and are presented in figures 3.8 and 3.9 respectively.

53

Open Impedance Mapping System

Start data collecting sequence

Switch SFB7 electrodes via

multiplexers to next tetrapolar set

SFB7 Impedance measurement

Rotate electrode orientation by 90o

Are all

electrode orientations

complete?

Are all

tetrapolar sets

complete?

Export raw impedance data to Matlab

Perform data analysis and

R0 calculations

Import R0 values to IMS program

Scale map ledgend for greatest

image contrast

Display impedance maps

No

Yes

Yes

No

Figure 3.7: Flow chart of the IMS operational sequence.



−1.5 −1 −0.5 0 0.5 1 1.5 2 2.5

−2

−1.5

−1

−0.5

0

0.5

Resistance Ohms

− R

eact

ance

Ohm

s

Figure 3.8: Measurement made with short circuited electrodes and no multi-plexer front end.

An analysis of variance (ANOVA) was performed on the data gathered for

resistance and reactance components to determine if a significant difference

was measured with and without the multiplexer front end. For analysis of

resistance the F statistic was 1.10, with a p-value of 0.30. Since this is greater

than 0.05 (or equivalently the observed F statistic is smaller than its critical

value of 3.86) it can be concluded that there is no significant difference between

the measured resistance with and without the multiplexers.

Analysis of the reactance produced a F statistic of 5.57, with a p-value

of 0.02. Since this is less than 0.05 (or equivalently the observed F statistic

is greater than its critical value of 3.86) it can be concluded that there is a

55

significant difference between the measured reactance with and without the

multiplexers.

−2 −1 0 1 2 3

−3.5

−3

−2.5

−2

−1.5

−1

−0.5

0

0.5

Resistance Ohms

− R

eact

ance

Ohm

s

Figure 3.9: Measurement made with short circuited electrodes through multi-plexer front end.

The result of no significant difference with resistance demonstrates the mul-

tiplexers Ron does not alter the resistive measurements. However significant

difference was found between the reactance values. This difference is due to

the close proximity of the multiplexers to each other and the electrode paths

taken on the PCB to the electrode array. This close proximity of the electrodes

needed to map small tissue sample has resulted in capacitive coupling of the

electrodes at high frequencies.



3.6.2 RRC Circuits

Various electrical circuits representative of biological tissue were used to test

the frequency response of the multiplexer front end. Figure 3.10 shows one

such RRC circuit consisting of a 100 Ω resistor in parallel with an in series

100 Ω resistor and 100 nF capacitor. The RRC circuits were measured with a

SFB7 in the range of 4 kHz to 1 MHz, with and without the multiplexer front

end attached. With the front end attached the RRC circuit was attached to

the first 4 multiplexer channels on the PCB electrode array board as shown

in figure 3.11. Buffer resistors of 100 Ω were used between the drive and

measurement electrodes. Typical Cole plots obtained with and without the

multiplexers are shown in figure 3.12.

100 Ω

100 nF100 Ω

Figure 3.10: RRC test circuit.

Measurements made without the multiplexer front end agreed within 1.5%

with the theoretical values for R0 and R∞

of 100 Ω and 50 Ω respectively.

This is not the case with the multiplexers attached, a second arc is produced

at frequencies higher than 90 kHz. This is due to the design of the printed

circuit boards manufactured for the multiplexers and electrode array. For a

small surface electrode array to be constructed the paths taken by the drive

57

R R

R

R C

Drive

Electrodes

Measurement

Electrodes

Figure 3.11: RRC test circuit attached to the electrode array. All resistorvalues (R) are 100 Ω and capacitor value (C) 100 nF.

and measurement electrodes must be very close. The close proximity of these

electrode paths results in a capacitive coupling, effectively ’shunting’ the RRC

test circuit at high frequencies.

Capacitive coupling of the electrodes may be modelled and taken into ac-

count during the data analysis, subsequently removing its effect. This will be

shown in following sections to be more difficult than anticipated and was not

the solution pursued. Alternatively higher frequencies (above 90 kHz) can be

removed from analysis. When this was done the resulting Cole plot agreed

within 1.5 % with theoretical R0 and R∞

values.



40 50 60 70 80 90

5

0

5

10

15

20

25

30

35

40

Ri = 48.2291 Ro = 99.6802

50 55 60 65 70 75 80 85 90 95

5

0

5

10

15

20

25

30

Ri = 49.9059 Centre = (74.2563,0.26315) Ro = 98.6066

Fc = 1.4379e008

R = 98.6066ohms S = 101.0467ohms C = 55439265446.6937uFF

igur

e 1

Resistance Ohms

Resistance Ohms

- R

eacta

nce O

hm

s-

Reacta

nce O

hm

s

(a)

(b)

90 kHz

Figure 3.12: Typical Cole plots for RRC test circuit, measured data is repre-sented by red circles and line of best fit by the broken blue line. (a) Cole plotobtained without the multiplexers (b) Cole plot obtained with the multiplex-ers.

59

3.6.3 Resistor Matrix

Mapping capabilities of the system were tested using an array of resistors con-

structed onto the electrode PCB, this resistor matrix is shown schematically

in figure 3.13. Measurements of a resistor matrix allowed testing of the mul-

tiplexers stepping sequence and the usable frequency range. The following

results are for a 1000 Ω resistor matrix and are typical of results also obtained

for 150, 330 and 560 Ω resistor matrices.

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

R

RR RR RR

RR

RR

RR

Figure 3.13: Resistor matrix constructed with 1000 Ω resistors.

A typical Cole plot and impedance map for measurements made with the

entire frequency spectrum (4 to 1000 kHz) are displayed in figures 3.14 and 3.15

respectively. The Cole plot no longer displays the second arc in the frequencies

above 90 kHz as previously seen with RRC circuits (figure 3.12b). This lack

of a second arc is due to the resistor matrix being non-complex (no reactive

component) as opposed to the RRC circuits, hence it is not the high frequency



arc (as seen in figure 3.12b) that is absent but the low frequency arc between

4 to 90 kHz.

The capacitive coupling of the electrode paths is still very evident in fig-

ure 3.14. Ideally all of the measured points should lie on the resistance axis

(zero reactance) at a single point. The measurements at high frequencies con-

tain a very small resistive component and a large reactive component. The

Cole plot obtained indicates that the majority of the current flow at these

frequencies is between the electrode PCB paths and not through the resistor

matrix. Capacitive coupling is typically overcome by shielding the electrodes,

however this was not possible due to size constraints.

Impedance mapping of a resistor matrix as shown in figure 3.15 also demon-

strates the effect of capacitive coupling in the map of R∞

. For a non-complex

electrical component, high frequency measurements should remain the same,

producing the same map as R0. Modelling the capacitive coupling effect and

taking this into account during the data analysis stage is a possible solution

for this problem. However with 64 independent measurements, each having a

unique PCB path through the multiplexers and onto the electrode array this

approach becomes time consuming and much larger than the scope of this

study. Limiting the frequency range used to less than 90 kHz will also not be

of use, as previously described with RRC circuits, since capacitive coupling

will still distort the measurements.

Variation in calculated values for R0 shown in figure 3.15a are not a result

of this capacitive coupling but size restrictions in the resistor matrix. If the

matrix was to be expanded the outer measurement will become closer to that

of the central values due to more resistors being in parallel.

61

-50 0 50 100 150

20

0

20

40

60

80

100

120

Ri = -57.0581 Ro = 152.8678

1 MHz

Resistance Ohms

-R

eact

ance

Ohm

s

4 kHz

Figure 3.14: Typical Cole plot for a measurement made with a resistor matrix,measured data is represented by red circles and line of best fit by the brokenblue line. This Cole plot was obtained from a matrix constructed of 1000 Ωresistors.

The lack of resistors becomes more evident when comparing the four smaller

maps of R0 obtained with different electrode orientations. The smaller maps

on the left have a higher resistance on the left and right sides due to more

current being forced to pass through the resistor between the measurement

electrodes. This current induces a larger voltage and resulting impedance

calculated. The same effect is seen with the maps on the right however their

electrode orientation produces regions of higher impedance on the upper and

lower sides.

The use of a resistor matrix has demonstrated that the electrode stepping



(a) (b)

Figure 3.15: Impedance map obtained from a 1000 Ω resistor matrix, this isrepresentative of a typical matrix result. (a) R0 map. (b) R

∞map.

sequence works and the results can be displayed as a surface map. However the

high frequency component of the measurement spectrum is still unreliable, but

the calculated Cole parameter R0 that relies on low frequency measurements

is still useable.

3.6.4 Biological Tissue

The impedance mapping system was tested using bovine blood as a biological

tissue. This was readily available and if needed could be obtained in large

volumes. Bovine blood fitted the RRC circuit shown in figure 3.10 with plasma

63

representing the resistive extracellular space and the red blood cells having a

capacitive membrane and resistive intracellular space.

Preliminary testing was performed using plasma extracted from bovine

blood. Plasma can be treated as a non-complex electrical component like a

resistor, since no red blood cells (or insignificant numbers) are present to act

as capacitors. BIS measurements should result in a single resistance value

(R0) with no reactance. Figure 3.16 displays a Cole plot obtained from an

impedance map measurement, and the same arc seen with the resistor matrix

is once again seen here due to capacitive coupling. The first arc present in the

RRC Cole plots is not present here because of the absence of red blood cells

to act as capacitance and intracellular resistance.

0 10 20 30 40 50

10

5

0

5

10

15

20

25

30

35

Ri = 5.1502 Resistance Ohms Ro = 53.8929

Rea

ctan

ce O

hms

1 MHz

4 kHz

Figure 3.16: Cole plot of pure plasma.



Impedance maps of R0 and R∞

are given in figure 3.17. Again the results

are similar to those for the resistor matrix, with R0 producing reliable values.

However the impedance map of R∞

produces values much less than expected

due to measurements being greatly affected by capacitive coupling.

(a) (b)

Figure 3.17: Impedance map of pure plasma. (a) R0 map. (b) R∞

map.

3.6.5 Testing Summary

Testing of the impedance mapping system on electrical circuits used to mimic

the electrical response of biological tissue, has shown that the multiplexer front

end is susceptible to producing erroneous results. Capacitive coupling of the

electrodes through the multiplexer and electrode array PCB was unavoidable

65

due to size constraints and hence close proximity of the electrode paths. How-

ever further advances in the multiplexer PCB design and construction may be

able to include shielding for the electrode paths, greatly reducing the effect of

capacitive coupling. Removal of the capacitance effect through modelling and

further data analysis would also prove difficult due to the multiple (64) indi-

vidual electrode paths. It is also not possible to just use the lower frequency

range below 90 kHz which was less effected in RRC circuits but clearly effected

when measurements over the entire array were made.

As an outcome of these results all further work conducted in this study will

concentrate on the R0 value. It is the least effected by capacitive coupling and

still provides significant change in measurable impedance, as will be demon-

strated in chapter 4. R0 also provides the greatest accuracy to extracellular

resistance compared to using a single frequency measurement (Cornish et al.,

1993). Measurements at lower frequencies are also where the majority of the

impedance change is seen between healthy and cancerous tissue. Brown et al.

(2000a) presented mean R0 values of 19.0 and 3.85 Ω m for normal squamous

epithelium and CIN 2/3 respectively. Whereas intracellular resistance, which

is determined from R∞

only has a range of 2.31 to 6.10 Ω m for normal squa-

mous epithelium and CIN 2/3. This larger range seen between extracellular

resistance will provide more accurate identification of suspected lesions.

Chapter 4

Bioimpedance Mapping -

Results and Discussion

The aim of this study was to gather impedance maps of freshly excised cervical

tissue. However as the project developed it became evident that this aim could

not be realised and an alternate approach to establishing the efficacy of the

bioimpedance mapping technique was necessary. Specific issues were identified

as problematic in the in-vitro assessment of excised tissue but which would not

be of concern in-vivo. Sample size of the excised tissue would not sufficiently

cover the electrode array or be of great enough thickness. In addition the

limited size of excised tissue would confine the current flow injected from drive

electrodes and result in erroneous impedance measurements, particularly on

the periphery of the sample.

For these reasons the efficacy of the bioimpedance mapping technique was

tested using bovine blood. The blood for each trial was collected from the

67

68CHAPTER 4. BIOIMPEDANCE MAPPING - RESULTS AND

DISCUSSION

same animal and treated with 70 mg/L of heparine to prevent coagulation.

Blood for each measurement was prepared in the same manner by allowing

it to cool to room temperature (22 degrees Celsius) and the red blood cells

separated via a centrifuge. The separated red blood cells and plasma could

then be mixed in appropriate proportions to obtain the required haematocrit

for testing. Samples were also collected and allowed to coagulate, these were

used to represent a high impedance tissue medium at R0 due to the small

extracellular space.

4.1 Effect of Tissue Sample Size

4.1.1 Electrode Array Coverage

FEA modelling indicated that it is important for the sample to sufficiently

cover the electrode array so the current flow is not confined as seen in figure 2.9.

Figure 4.1 shows an example of insufficient coverage in the upper right corner.

The result of this inadequate coverage is that a smaller volume of conductive

medium is available for the current to flow through, increasing the measured

impedance in the upper right corner in relation to the rest of the impedance

map. The increase in impedance (Z) is due to it being inversely proportion to

the conductor volume (V) as given in equation 4.1 (ρ = resistivity and L =

conductor length).

Z =ρL2

V(4.1)

69

Figure 4.1: Inadequate electrode coverage with bovine blood in the upper rightcorner.

Figure 4.2 displays the impedance map produced when the volume of

plasma in the upper right corner is reduced. The thickness of the sample

was 4 mm. The effect of varying the sample thickness is considered in the

following section 4.1.2. Here the dam that is used to contain the bovine blood

in place on the electrode array has been brought closer to the upper right

corner of the electrode array, reducing the amount of plasma overlap on the

electrodes. The increase in measured impedance, with respect to the rest of

the map, in this corner is clearly seen.

If biopsy samples were to be used, they must be of a sufficient size to

eliminate these impedance distortions. This would not have been feasible in

a clinical trial as the typical sample taken at biopsy is smaller than the area

of the electrode array. With a biopsy sample not covering the electrode ar-


DISCUSSION

Figure 4.2: Impedance map with a higher impedance in the upper right cornerdue to insufficient electrode coverage.

71

ray, tetrapolar electrode sets on the boundary of the sample will have partial

coverage resulting in heavily distorted impedance measurements.

4.1.2 The Effect of Sample Thickness

FEA modelling showed that the sensitivity decreased rapidly with depth and

that the thickness of the sample must be at least 5 to 6 mm (figure 2.23).

This was tested experimentally by confining the plasma sample with a glass

cover slip attached to a micrometer which could be used to alter the sample

thickness. Use of the glass plate and micrometer permitted a sample of known

thickness to be measured. Figure 4.3 shows the average R0 measurement made

for plasma at thicknesses of 0.25, 0.50, 1.00, 2.00, 3.00, 4.00 and 5.00 mm.

With an increase in plasma thickness the R0 value decreases substantially

to a constant value at a thickness of 4 to 5 mm. This is consistent with

the thickness indicated by the FEA modelling of 5 to 6 mm. Thus although

the thickness of the cervix is greater than 6 mm the bioimpedance technique

samples the surface layer where any precancerous changes occur. As a result

of this observation all following impedance maps were taken with a minimum

sample thickness of 4 mm.

It is evident that virtual biopsy of small tissue sizes will be prone to prob-

lems of confined current flow. If impedance mapping is to be continued for

use on excised tissue then a sample size sufficient to more than cover the array

will be needed to avoid distorted maps at the tissue boundaries and edges of

the electrode array. The thickness of the sample would also need to be at least

4-5 mm. As mentioned above, biopsy samples of cervical tissue are unlikely to


DISCUSSION

0 1 2 3 4 540

50

60

70

80

90

100

Thickness (mm)

Mean

R 0

Figure 4.3: Mean R0 of plasma for various sample thicknesses.

73

meet these requirements. However further development of the instrumentation

such as miniaturisation and wireless communication between a personal com-

puter and recording device would allow for measurements to be made in-situ,

minimising the distortion of impedance measurements as a result of insufficient

sample size.

4.2 Homogeneous Haematocrit Impedance Map-

ping

The sensitivity of the impedance mapping system (as described in chapter 3)

to changes in impedance values (R0) was studied using blood samples with a

range of haematocrit values. A significant impedance range would ensure an

impedance map with differing haematocrits would show defined boundaries.

Impedance maps were obtained using homogeneous blood samples with a

range of prepared haematocrit values (0, 20, 40, 60, 80%). A haematocrit of

0% represents pure plasma. The individual haematocrit samples were placed

on the electrode array one at a time, with a volume large enough to completely

cover the electrode array and 4 mm in thickness.

Figure 4.4 shows the resultant impedance maps for the samples. It can be

seen that the impedance maps display constant R0 values for individual maps

(CV = 3 %). This confirms the results for testing the Ron contribution in

section 3.6.1. If Ron was to contribute to the measured impedance it would

result in variations of R0 values throughout the map.


DISCUSSION

0

250

0%

20%

40%

60%

80%

Figure 4.4: Homogeneous haematocrit impedance maps. The haematocrit isgiven as a percentage and the colour legend displays the R0 value in ohms.

75

0 10 20 30 40 50 60 70 8040

60

80

100

120

140

160

180

200

Haematocrit %

Me

an

R0

R2 = 0.97

Figure 4.5: Mean R0 for homogeneous impedance maps.


DISCUSSION

The constant R0 values also show that the inactive electrodes are not inter-

acting with the active measurement by creating a short circuit. If this was the

case a different impedance would have been measured in the outer 12 regions

to that of the 4 inner regions.

A plot of mean R0 for each impedance map against haematocrit (figure 4.5)

shows a large increase in impedance with haematocrit. The plot follows an

exponential trend as expected since the R0 value of a sample with haematocrit

of 100 % would approach infinity due to the very small extracellular space. The

range of haematocrit values has also shown to have a significant and measurable

change in R0. This was vital if impedance maps were to be measured with two

or more volumes of blood with differing haematocrits.

4.3 Impedance Maps

4.3.1 Plasma with Introduced Red Blood Cells

Impedance maps of samples of non-homogeneous haematocrits were measured

in the same manner as used for homogeneous haematocrits to investigate if

a region of different haematocrit in an otherwise homogeneous sample could

be identified. The electrode array was covered with plasma (0% haematocrit)

and regions of higher haematocrit introduced onto the electrode array via a

hypodermic syringe. This was not a stable arrangement as the introduced

RBC diffused through the plasma. Figure 4.6 demonstrates a region of red

blood cells (100% haemotcrit) introduced in the lower left corner of a plasma

sample and the associated diffusion through the plasma.

77

Figure 4.6: Region of red blood cells introduced to plasma. Lower left darkarea is where the cells were injected and the light grey is the area of visiblediffusion.

Cells were introduced via a hypodermic syringe onto the lower left electrode

as a region of higher impedance and to mimic a lesion on the surface of healthy

tissue. Impedance maps were obtained immediately after the red blood cells

were introduced and examples of typical results obtained from this process

are shown in figure 4.7 and 4.8. More red blood cells were introduced for the

impedance map shown in figure 4.8.

A clear distinction between the impedance of the upper right corner (pure

plasma) can be seen compared with the lower left corner (introduced red

blood cells). In both impedance maps the region of introduced cells can be

clearly identified by regions of high impedance. Figure 4.8 indicates a higher

impedance in the lower left as a result of more introduced RBC compared to


DISCUSSION

Figure 4.7: Impedance map of plasma with red blood cells introduced in lowerleft corner.

79

figure 4.7. However the boundaries are not clear due to the dispersion of red

blood cells into the plasma. The dispersion is identified by the gradual change

in impedance between the upper right and lower left corners. A schematic of

dispersion is shown in figure 4.6.

Figure 4.8: Impedance map of plasma with a larger volume of red blood cellsintroduced in lower left corner.


DISCUSSION

Red Blood Cell Dispersion

The dispersion of red blood cells throughout the plasma was expected to affect

the impedance measurements by continuously changing the impedance mea-

sured in each region. To measure this effect red blood cells were introduced in

the lower left corner of the plasma and measurements made at 2 minute inter-

vals (the time taken for one measurement). Figure 4.9 shows the 5 maps taken

over a ten minute period and the resulting decrease in measured impedance

over time in the lower left corner.

The continual change in impedance is not desirable since a single measure-

ment can take 2 minutes and impedance values will change during this single

measurement. The dispersion effect would not be seen in a biopsy and in-vivo

measurements.

4.3.2 Plasma with Introduced Red Blood Cell Clot

To stop the dispersion of red blood cells a clot was introduced to the plasma

in place of the RBC. This should result in regions of higher impedance and

more defined boundaries, with no gradual change in impedance from the region

of the clot to regions of plasma. Figures 4.10, 4.11 and 4.12 display typical

impedance maps with clots introduced in various regions on the electrode array.

The maps measured show clear impedance changes at the boundaries of the

red blood cell clots due to minimal dispersion of RBC. With defined boundaries

it is now possible to accurately mimic a surface lesion and precisely identify

the size of a lesion and monitor its growth or change in shape over time.

81

(a) (b)

(c) (d)

(e)

Figure 4.9: Dispersion of introduced red blood cells into plasma over 2 minuteintervals. (a) Measurement made immediately after the introduction of RBC.(b) Measurement made at 2 minutes. (c) Measurement made at 4 minutes.(d) Measurement made at 6 minutes. (e) Measurement made at 8 minutes.


DISCUSSION

Figure 4.10: Impedance map of plasma with introduced red blood cell clotcovering central electrode.

83

Figure 4.11: Impedance map of plasma with introduced red blood cell clotcovering the 4 electrodes associated with the region in the middle lower right.


DISCUSSION

Figure 4.12: Impedance map of plasma with introduced red blood cell clotcovering the 2 lower right regions.

85

4.4 Experimental Comparison with Modelled

Sensitivity Field

4.4.1 Anomalous Measurements

Evidence of the anomaly found in FEA modelling (see section 2.2.2) and as a

result of the tetrapolar configuration’s complex sensitivity field is also shown in

the experimental impedance maps. Figure 4.13 presents results obtained from

the introduction of red blood cells to plasma and with subsequent dispersion

as displayed in figure 4.6. The map shows the red blood cell density and re-

sulting impedance is higher between the 2 lower electrodes associated with the

4 electrodes used to measure the region identified by the arrows in figure 4.13.

The sensitivity region between these 2 electrodes is positive for the smaller

maps on the right, resulting in a increased measured impedance if a higher

impedance medium is here. This increase in impedance is clearly seen in the

two maps on the right and indicated by arrows. The two maps on the left show

a decrease in impedance because the region between the lower electrode pair

is of negative sensitivity, resulting in a decreased measured impedance when a

higher impedance medium is located in the region.

This measured anomaly was seen in nearly every impedance map recorded

and agrees with the sensitivity fields modelled with FEA. However it does not

alter the averaged impedance map since the R0 value in this region of the

larger map (average impedance) is not anomalous i.e. the R0 value follows the

decreasing impedance gradient from the lower left corner to the upper right.


DISCUSSION

(1) (2)

(3) (4)

Figure 4.13: Anomalous result due to positive and negative sensitivity fields.Anomalous result is identified by arrows.

87

4.4.2 Reciprocal Electrodes

The reciprocal nature of the tetrapolar electrode configuration as indicated

by FEA modelling is clearly seen in figure 4.14. The two smaller electrode

orientation maps on the left are the same as are the pair on the right. The only

differences within the pairs are that the drive and measurement electrodes have

been interchanged. This can be explained by consideration of the symmetrical

nature of the sensitivity fields as shown in section 2.3.

The reciprocal nature of the tetrapolar electrode configuration suggests

that the impedance map might be obtained using only two electrode orienta-

tions rather than the four currently proposed. The number of required mea-

surements could therefore be halved. However the uncertainty in the measured

impedance is also important.

The uncertainty was tested using an impedance map obtained for a pure

plasma sample (figure 4.15). The average of all four is the current proposed

approach (1+2+3+4), and the average of new electrode orientations usable

are 1+2, 3+4, 1+4 and 3+2. The mean and variance of R0 for these averages

are presented in table 4.1.

Combined Maps Mean Variance

1,2,3,4 49.59 1.961,2 49.57 2.063,4 49.61 1.911,4 49.61 1.673,2 49.58 2.35

Table 4.1: R0 mean and variance for various electrode orientation combina-tions.

Analysis of variance (ANOVA) for R0 produced a F statistic of 0.002, with


DISCUSSION

Figure 4.14: Reciprocal nature of the tetrapolar electrode configuration. Thereciprocal pairs are grouped

89

21

3 4Figure 4.15: Impedance map of plasma used to demonstrate the reciprocalnature of the tetrapolar electrode configuration.


DISCUSSION

a p-value of 1.0. Since the p-value is greater than 0.05 (or equivalently the ob-

served F statistic is smaller than its critical value of 2.494) it can be concluded

that there is no significant difference between the measured mean R0 values

using the different pairs of electrode orientations and the original approach of

using all 4 orientations. Similar results were also obtained for other impedance

maps analysed.

Since the averages were shown to be not statistically different it can be

accepted that only 2 impedance measurements are required at each region.

These measurements can also be of any electrode orientation as long as they

are 90o to each other. Any anomalies present will also still be removed by

averaging since there is no difference between averaging 4 measurements where

there are infact only 2 unique measurements, or just averaging these 2 unique

measurements.

Efficient Electrode Stepping

As shown above there is no advantage in reproducibility in averaging four

rather than two maps. It is now possible to take advantage of the reciprocal

nature of the tetrapolar configuration and utilise a more efficient electrode

stepping sequence. Figure 4.16 displays an alternative method for stepping

through the electrode array, here the number of electrodes can be double that

of the available multiplexer channels.

Using this presented sequence a 25 electrode array may be constructed

with only 12 channel multiplexers. Electrodes white (measurement) and black

(drive) will require only 12 channels each, while the other blue measurement

91

(1) (2)

(3) (4)

(5) (6)

(7) (8)

Figure 4.16: New proposed electrode stepping sequence 1-8. Red/black andwhite/blue represent drive and measurement electrodes respectively.


DISCUSSION

electrode requires 9 channels and the red even less with only 4 channels. This

stepping sequence was applied to an impedance map and is show in figure 4.17

along with the map obtained using an average of 4 measurements.

Comparison of the impedance maps show only minor changes in R0 values,

with some differences due to rounding in the data analysis. However the region

of higher impedance in the lower left corner is still clearly identifiable.

4.5 Lesion Boundary Identification

The anomaly confirmed in section 4.4.1 is only found in regions where a

medium of different impedance is located under the electrodes. A possible

method for detection of lesion boundaries is now available since 2 measure-

ments using orthogonal electrode orientations, at a region of non-homogeneity,

will produce different measured impedances. Whereas a region of homogeneity

will produce the same measurements. The impedance map of plasma with in-

troduced RBC shown in figure 4.13 was used for the following example. Smaller

maps 1 & 3 have been averaged , the same process was used for maps 2 & 4.

The difference between these 2 maps is presented in figure 4.18.

A lesion’s boundary can now be identified in the resultant map as a region

with a non-zero R0 value. For example a boundary map of homogeneous

plasma will result in values close to zero since there is no difference in the

medium under the electrodes. However in figure 4.18 it can be seen in the

region of dispersed blood there are high values due to different haematocrits

being located under each of these electrode sets. In the lower left a value of

93

(a)

(b)

Figure 4.17: Comparison of (a) new electrode stepping sequence with (b)presently used.


DISCUSSION

Figure 4.18: Boundary identification via anomalies.

95

zero is seen due to a high but homogeneous haematocrit sample being located

under the electrode set. This is the location of the injected RBC. The upper

right region of homogeneity is also close to zero due to the sample under the

electrode sets being homogeneous plasma, the RBC have not dispersed into

this region. This has identified a method for objectively determining lesion

boundaries.

4.6 Summary

Virtual biopsy of small tissue sizes will be prone to problems of confined cur-

rent flow. If impedance mapping is to be continued for use on excised tissue

then minimum size requirement will be needed to avoid distorted maps at the

tissue boundaries and edges of the electrode array. Further development of the

instrumentation and electrodes would allow for measurements to be made in-

situ, eliminating all effects and distorted impedance measurements as a result

of insufficient sample size.

The use of bovine blood as a substitute for freshly excised tissue proved to

be acceptable. A significant change in impedance was seen with a change in

haematocrit and these impedance changes were measurable in the mapping.

The advantage of using a clot to confine the regions of high impedance and

provide distinct boundaries also proved useful.

Anomalous results indicated during modelling of the tetrapolar configu-

rations sensitivity fields in chapter 2 are also present in the experimental

impedance maps. However they are only found when lesion boundaries or


DISCUSSION

changes in medium impedance are located underneath the electrodes. These

anomalies did not appear to alter the resultant impedance map once averaged.

This negated the need to remove and discard this measurement as previously

expected. A surprising side effect of the anomaly is the ability to now detect

boundaries of lesions or changes in impedance. This allows a non-subjective

method for determining lesion size and possible biopsy margins.

The reciprocal nature of the electrodes shown by modelling of the sen-

sitivity fields agree with the experimental impedance maps measured. This

also provides a method for increasing the number of electrodes in the array

while still using multiplexers with the same number of channels. Reducing the

number of measurements required at each region will also minimise the time

taken to acquire an impedance map. However, despite halving the number of

measurements from four to two at each region statistically there is no decrease

in the measurement accuracy.

Chapter 5

Conclusion

Virtual biopsy by bioimpedance spectroscopy is a relatively new technique.

This study investigated the sensitivity fields of the commonly used tetrapolar

electrode configuration and expanded it to consider the sensitivity of a 25

electrode array. The use of an array allowed further characterisation of a lesion

of interest by identifying spatial information, importantly lesion boundaries.

This work into mapping the surface impedance of a lesion has not previously

been undertaken as other investigations of virtual biopsy have used single point

measurements on a region of interest, with no additional information gathered

about the lesion dimensions.

Finite element analysis (FEA) was used to study the sensitivity for the

measurement of impedance of tissue using the tetrapolar electrode configura-

tion with simulated lesions located in healthy tissue. As expected, a decrease

in measured impedance resulted when a lesion was introduced, except for the

case of a lesion located between a drive and measurement electrode pair. In

97

98 CHAPTER 5. CONCLUSION

this latter case, the measured impedance was found to initially increase when

small lesions were present and then decrease as the lesion size increased. This

“delayed” decrease in measured impedance could lead to a false negative re-

sult, leaving the lesion undetected and subsequently not identified for further

treatment.

Modelling of sensitivity fields provided insight of the distribution of areas

of positive and negative sensitivity. These fields were found to agree with and

explain the previous anomalous results indicated by the modelled lesions. The

rapid decrease in sensitivity with depth confirmed the efficacy of using the

tetrapolar electrode configuration for surface measurements. Results showed

that the magnitude decreased to near zero at a depth of 5 mm, confirming the

measurement sensitivity to the surface layers.

A multi-electrode array showed surprising results in that the inactive elec-

trodes surrounding the active electrodes had little effect on the measured

impedance. Modelled sensitivity fields indicated additional regions of pos-

itive sensitivity between the active and inactive electrodes. However these

had a magnitude 10 times less than those found within the active electrodes,

explaining their lack of effect on measured impedance. They were shown, ex-

perimentally, to not affect the measured impedance in maps of homogeneous

haematocrit samples

Instrumentation has been designed and constructed allowing for the use of

an electrode array with a commercially available impedance measuring device.

The multiplexer front end developed in this project allows for the drive and

measurement electrodes of the commercial device to be switched to any of the

25 electrodes in the array. The front end is also designed to be used on a

99

desktop personal computer without the need for any specialised equipment.

A printed circuit board based electrode array was developed and imple-

mented. A PCB based electrode arrays also makes manufacturing of arrays

readily available and disposable, but autoclavable for repeated use if desired.

Automation of the multiplexing front end and data analysis process was

performed via Matlab and a Visual Basic program. The Visual Basic program

provided an objective display of the measured impedance map for the user.

Testing of the multiplexer front end showed that it suffers from capaci-

tive coupling. This did not allow for accurate impedance measurements at

high frequencies and the use of R∞

as an analysis parameter. However R0

was unaffected by the capacitive coupling and still usable to provide detailed

impedance maps with high contrast between regions of different impedance.

A small tissue sample size was determined to be a problem and trials with

freshly excised cervical tissue would not be possible since adequate coverage

of the electrode array by the sample is required as well as adequate sample

thickness. It was shown that if these size conditions were not met then confined

current flow throughout the tissue medium would result in distorted impedance

measurements.

A substitute tissue medium was found in bovine blood and mapping of

homogeneous haematocrit demonstrated a large range of impedance values

that would be suitable for impedance mapping. The Impedance Mapping

System (IMS) was shown to measure these regions of impedance change when

red blood cells and clots were introduced to plasma, displaying them clearly

in the maps produced. Changes in impedance at the boundaries resulted in

100 CHAPTER 5. CONCLUSION

anomalous impedance measurement as modelled in FEA. It was thought these

anomalous results would need to be removed from further analysis to prevent

altering the final outcome. However averaging of the maps obtained with

electrodes rotated by 90o showed no evidence of anomalous effects. A method

for objectively identifying lesion boundaries utilising these anomalous results

was developed.

Taking advantage of the reciprocal nature of the tetrapolar electrode config-

uration a more efficient electrode stepping sequence was developed. Without

changes to the present instrumentation the number of measurements, com-

pared to that initially considered, may be halved without any detrimental

effects to the measurement quality, resulting in faster measurement times and

data analysis.

In summary, an impedance mapping system has been modelled, designed

and developed for tissue characterisation by bioimpedance measurements. The

technique has been shown experimentally to be able to detect regions of differ-

ent impedance and is in agreement with the finite element analysis performed.

Further development of the IMS will allow progressive monitoring of suspect

lesions in-vivo and better identification of their spatial distribution for biopsy.

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