optical coherence tomography for optical biopsy of

Optical coherence tomography for �“optical biopsy�” of axillary lymph nodes

involved in breast cancer metastasis

Loretta Scolaro BE(Hons), BSc

This thesis is presented for the degree of

Doctor of Philosophy

of The University of Western Australia

School of Electrical, Electronic & Computer Engineering.

January 2014

iii

Abstract One of the pathways for the spread of metastatic breast cancer throughout the body is the

lymphatic system. Accurate determination of the presence and extent of metastatic deposits in

lymph nodes of the axilla (staging) is critical to the management of the disease. Currently,

lymph nodes are assessed by microscopic examination after surgical excision. However, this

can mean removal of healthy lymph nodes without metastatic disease, and may result in

unnecessary lymphoedema; that is, swelling due to accumulation of tissue fluid as a result of

disruption to normal lymphatic pathway function. This thesis presents an investigation into a

possible optical method of lymph node assessment based on optical coherence tomography

(OCT) that has the potential to provide in situ “optical biopsy” of lymph node involvement.

In situ biopsy could reduce the rate of unnecessary lymphoedema and improve efficiency of

node assessment.

In this thesis, an investigation of the baseline appearance of the morphology of human

axillary lymph nodes imaged using OCT is first presented. We assess OCT images of excised

human axillary lymph nodes; both healthy nodes and nodes involved in metastatic spread of

breast cancer are imaged using a benchtop OCT system. We compare OCT images with the

structural and cellular composition identified using the gold standard technique of

histopathology. The results identify a higher OCT signal from metastatic deposits and show

that OCT can successfully image the micro-architecture in lymph nodes.

It was evident from this study, however, that the diagnostic capability of OCT requires

improvement. The native contrast of intensity-based OCT is not, in itself, sufficient to

differentiate metastatic deposits in involved lymph nodes, primarily because interpretation of

such contrast is ambiguous and subject to artefacts. One of the most problematic artefacts

occurs as a result of the variable attenuation of light in tissue. In this thesis, we therefore

present a method for processing the OCT data to overcome these ambiguities. The method

exploits the attenuation of light in tissue to provide a quantitative measure of contrast. It

utilises a single-scattering model to describe the OCT reflectance from tissue and extract

optical attenuation coefficients. The attenuation coefficient is an optical property of tissue

that is determined by its unique cellular and structural composition and is, therefore, expected

to vary between healthy and malignant tissue. We demonstrate improved differentiation of

healthy and abnormal lymph node tissue types based on measurement of this coefficient and

iv

use a visual representation to interpret the measured values by mapping them into en face

parametric images.

A second improvement we propose for enabling OCT optical biopsy is that of advancing

the technology into a needle suitable for in situ imaging in a clinical setting. To this end, we

present the design, fabrication and optical characterisation of two fibre-optic needle probes

that we use for in situ OCT needle imaging of breast tumour tissue and lymph nodes. We

demonstrate that the second of these needle probes achieves the highest sensitivity for needle

probes reported to date. Furthermore, we present the use of this needle to measure the optical

attenuation coefficients of excised, unprocessed lymph nodes in a proof of principle study.

This study demonstrates the improved capability of parametric OCT for differentiating

malignant and healthy lymph nodes. This thesis supports the possibility of OCT as an optical

biopsy tool for in situ axillary lymph node diagnosis.

v

Contents Abstract iii

Contents v

List of figures ix

List of tables xiii

Acknowledgements xv

Statement of Contribution xvii

List of publications xxi

List of Acronyms xxv

Chapter 1 Introduction 1

1.1 Thesis motivation ................................................................................................................ 6

1.2 Structure of this thesis ........................................................................................................ 6

Chapter 2 Background 9

2.1 Breast cancer and lymph node metastasis ....................................................................... 9

2.1.1 Disease progression ...................................................................................................... 10

2.1.2 Staging and treatment.................................................................................................. 11

2.1.3 Morphology of lymph nodes...................................................................................... 11

2.1.4 Standard histopathological evaluation..................................................................... 12

2.1.5 Side-effects of histopathological evaluation............................................................ 13

2.1.6 Sentinel lymph node biopsy ....................................................................................... 13

2.1.7 Motivation for a new assessment technique ........................................................... 15

2.2 Alternative techniques for in situ assessment of lymph node status ......................... 15

2.2.1 Pre-operative needle biopsy........................................................................................ 16

2.2.2 Pre-operative imaging .................................................................................................. 16

2.2.3 Novel techniques in optical diagnosis ...................................................................... 18

2.3 Background of optical coherence tomography.............................................................. 19

2.3.1 Basic principles of OCT.............................................................................................. 22

2.3.2 OCT systems used in this thesis................................................................................ 24

Chapter 3 OCT imaging of axillary lymph nodes 29

3.1 Related work ......................................................................................................................... 30

3.2 Optical scattering and lymph node tissue types ............................................................ 31

3.2.1 Origins of scattering in tissue..................................................................................... 31

vi

3.2.2 Histology and scattering predictions for healthy lymph node tissues............... 35

3.2.3 Histopathology and scattering predictions for metastatic carcinoma in

lymph nodes................................................................................................................... 40

3.3 Imaging of human lymph nodes using optical coherence tomography.................... 44

3.3.1 Introduction .................................................................................................................. 44

3.3.2 Materials and methods ................................................................................................ 46

3.3.3 Results ............................................................................................................................. 47

3.3.4 Discussion ...................................................................................................................... 51

3.3.5 Grant Support ............................................................................................................... 52

3.3.6 References for Section 3.3........................................................................................... 53

3.4 Results of intensity-based OCT for lymph node diagnosis ........................................ 55

3.4.1 Problems of intensity based differentiation............................................................ 56

3.4.2 Conclusions ................................................................................................................... 59

Chapter 4 OCT for optical property measurement of lymph nodes 61

4.1 Optical properties background ......................................................................................... 62

4.1.1 Definitions of OCT-related optical properties...................................................... 63

4.1.2 Assumptions for optical property measurement in tissue ................................... 67

4.2 OCT for optical property measurement ........................................................................ 68

4.2.1 The single-scattering model of OCT ....................................................................... 80

4.3 Parametric imaging of the attenuation coefficient in human axillary lymph

nodes ...................................................................................................................................... 81

4.3.1 Experimental methods................................................................................................. 82

4.3.2 Relative attenuation based parametric OCT ......................................................... 83

4.3.3 Quantitative attenuation coefficient based parametric OCT ............................ 87

4.4 Discussion and conclusions ............................................................................................... 100

Chapter 5 Needle-based OCT 101

5.1 OCT imaging endoscopes and needles in the literature.............................................. 102

5.1.1 Side-imaging probes ..................................................................................................... 102

5.1.2 Forward imaging probes.............................................................................................. 105

5.2 Imaging of breast cancer with optical coherence tomography needle probes......... 109

5.2.1 Introduction .................................................................................................................. 109


5.2.3 Results ............................................................................................................................. 119

vii

5.2.4 Discussion ...................................................................................................................... 121

5.2.5 Conclusion..................................................................................................................... 122

5.2.6 References for Section 5.2........................................................................................... 123

5.3 Discussion: towards an improved needle probe design ............................................... 127

5.4 High-sensitivity anastigmatic imaging needle for optical coherence tomography 129

5.4.1 Introduction .................................................................................................................. 129


5.4.3 Results ............................................................................................................................. 133

5.4.4 References for section 5.4 ........................................................................................... 136

5.5 Optical property measurement of lymph nodes using an OCT needle ................... 138

5.5.1 A portable OCT needle imaging system ................................................................. 138

5.5.2 Sample preparation and imaging protocol .............................................................. 140

5.5.3 Data processing ............................................................................................................. 141

5.5.4 Notes on the presentation and interpretation of results...................................... 142

5.5.5 Results ............................................................................................................................. 143

5.5.6 Discussion ...................................................................................................................... 147

5.5.7 Conclusions ................................................................................................................... 148

Chapter 6 Conclusions 149

6.1 Going forward – where to from here? ............................................................................. 151

6.2 Study limitations and future work ................................................................................... 153

6.3 Final remarks......................................................................................................................... 155

A Needle gauge chart 157

B OCT system calibration 159

C Parameters for TDOCT attenuation coefficient extraction 165

Bibliography 169

ix

List of figures Figure 1-1 Bar chart of Scopus search results of OCT biological applications over the

past ten years............................................................................................................... 4

Figure 2-1 Illustration of breast cancer and its metastasis to lymph nodes. ..................... 10

Figure 2-2 The structure of a typical lymph node.................................................................. 12

Figure 2-3 Severe lymphoedema in the right arm of a patient after undergoing

ALND for breast cancer staging............................................................................. 13

Figure 2-4 Lymphoscintigram transmission images taken prior to sentinel biopsy. ...... 14

Figure 2-5 Ultrasound images of a healthy and abnormal lymph node. ........................... 18

Figure 2-6 Diagram of the diagnostic window in tissue showing the spectral

absorption for a range of tissue chromophores................................................... 20

Figure 2-7 Resolution and imaging depth of OCT compared with other imaging

modalities. ................................................................................................................... 21

Figure 2-8 The OCT sample optics determine transverse resolution and axial depth

of field. ......................................................................................................................... 23

Figure 2-9 Schematic diagram of the fibre-based TDOCT system. .................................. 25

Figure 2-10 Schematic diagram of the Thorlabs SSOCT system. ........................................ 26

Figure 2-11 Schematic diagram of the custom-built portable SSOCT system.................. 28

Figure 3-1 Components of a soft tissue that cause spatial refractive index variations. .. 32

Figure 3-2 The typical organisation of a cell. .......................................................................... 33

Figure 3-3 Relative backscattering intensity is determined by cellular composition...... 35

Figure 3-4 Micrograph of the typical morphology of a healthy lymph node. .................. 36

Figure 3-5 Micrographs of the composition of lymph node cortical follicles. ................. 37

Figure 3-6 Micrographs of the composition of lymph node paracortex. .......................... 38

Figure 3-7 Electron microscope image of the medullary area of a lymph node. .............. 39

Figure 3-8 Electron microscope image of the capsule and marginal sinus of a lymph

node.............................................................................................................................. 39

Figure 3-9 Medium magnification micrograph of mature adipose tissue and high

magnification micrograph of three vessels. .......................................................... 40

Figure 3-10 Micrograph showing neoplastic cells of metastatic ductal carcinoma. .......... 41

Figure 3-11 Micrographs showing patterns of metastatic growth. ....................................... 42

Figure 3-12 H&E and OCT images of lymph node 1 and lymph node 2. ......................... 48

x

Figure 3-13 H&E and OCT images of lymph node 3 showing perinodal blood vessel. .. 49

Figure 3-14 H&E and OCT images of diffuse metastatic cancer in lymph node 4 and

lymph node 5.............................................................................................................. 50

Figure 3-15 H&E and OCT images of metastatic deposits in lymph node 6 and lymph

node 7........................................................................................................................... 51

Figure 3-16 Visualisation of three-dimensional OCT image of a lymph node showing

streak artefacts............................................................................................................ 57

Figure 3-17 TDOCT images of lymph nodes demonstrating shadow artefacts. .............. 58

Figure 3-18 TDOCT en face and YZ images of a lymph node demonstrating a gap

artefact. ........................................................................................................................ 58

Figure 3-19 TDOCT images of a malignant lymph node demonstrating artefacts due

to attenuation............................................................................................................. 59

Figure 4-1 Example of a scattering phase function generated using Mie theory. ............ 64

Figure 4-2 Absorption and scattering (i.e., attenuation) coefficients for various tissue

constituents across a broad range of wavelengths............................................... 66

Figure 4-3 Averaged A-scans from two phantom objects..................................................... 85

Figure 4-4 Involved (malignant) human axillary lymph node with diffuse

involvement of the node tissue............................................................................... 85

Figure 4-5 Involved (malignant) human axillary lymph node with diffuse

involvement of the node tissue............................................................................... 86

Figure 4-6 OCT cross-sectional images of a representative microsphere correction

phantom highlighting the confocal and scanning modulations...................... 89

Figure 4-7 Measured attenuation coefficients for a healthy lymph node. ........................ 91

Figure 4-8 Measured attenuation coefficients for a healthy lymph node with

prominent cortex and fatty hilum. ........................................................................ 93

Figure 4-9 Measured attenuation coefficients for a reactive lymph node with an area

of necrotic tissue containing dystrophic calcifications and a loss of normal

tissue architecture...................................................................................................... 95

Figure 5-1 Illustration of the typical scanning configuration for a side-imaging OCT

endoscopic probe....................................................................................................... 103

Figure 5-2 Designs and scanning mechanisms of side-viewing OCT probes................... 104

Figure 5-3 Designs and scanning mechanisms of forward-viewing OCT probes. .......... 106

Figure 5-4 Schematic of our OCT needle probe.................................................................... 112

xi

Figure 5-5 Results of the simulation of the Gaussian beam propagation in our GRIN

fibre probes. ................................................................................................................ 113

Figure 5-6 Schematic of the 836 nm SDOCT system used for the needle imaging

experiments................................................................................................................. 114

Figure 5-7 Measuring the absolute value of the reference reflectivity for the OCT

needle probe................................................................................................................ 116

Figure 5-8 OCT needle probe image of a tumour margin. .................................................. 119

Figure 5-9 OCT needle probe images of two lymph nodes. ................................................ 121

Figure 5-10 Schematic of our capillary encased OCT needle probe. ................................... 130

Figure 5-11 Illustration of the anastigmatic design concept. ................................................. 132

Figure 5-12 WD and spot-size ratios of our design goal and our fabricated probe........... 133

Figure 5-13 Measured and simulated beam profiles for our fabricated probe in air and

in water. ....................................................................................................................... 134

Figure 5-14 OCT needle A-scan of a water/glass interface.................................................... 135

Figure 5-15 Radial OCT needle image of the surface of a finger and thumb..................... 135

Figure 5-16 Radial and planar view of the 3-D OCT needle image of a cucumber. ......... 136

Figure 5-17 Noise characteristics of the portable 1310 nm SSOCT system. ..................... 139

Figure 5-18 Photo of the portable SSOCT system. ................................................................. 140

Figure 5-19 Photo of the needle probe connected to the rotation and translation

motors and the sample stage with sample holder. .............................................. 141

Figure 5-20 Needle OCT imaging and optical property measurement of a healthy

lymph node. ................................................................................................................ 144

Figure 5-21 Needle OCT imaging and optical property measurement of a malignant

lymph node. ................................................................................................................ 146

Figure 5-22 Needle OCT imaging and optical property measurement of a section of

breast tumour. ............................................................................................................ 147

Figure B-1 Sample arm setup for calibration of the TDOCT system................................ 161

Figure B-2 Calibration curve for the TDOCT system. ........................................................ 162

Figure B-3 Calibration curve for the custom-built needle SSOCT system. ..................... 163

Figure C-1 OCT scan of a microsphere sample used for selecting processing

parameters and verifying the procedure. .............................................................. 166

Figure C-2 Microsphere sample (corrected) reflectance profiles for different

averaging window sizes and axial fitting ranges. ................................................. 166

xiii

List of tables Table 2-1 Theoretical and measured performance parameters for the 1310 nm

TDOCT system. ....................................................................................................... 25

Table 2-2 Theoretical and measured performance parameters for the 1325 nm

Thorlabs SSOCT system......................................................................................... 27

Table 2-3 Theoretical and measured performance parameters for the 1310 nm

custom-built needle SSOCT system. .................................................................... 28

Table 3-1 Predictions of relative scattering intensity for axillary lymph nodes imaged

with OCT. .................................................................................................................. 43

Table 3-2 Resulting relative OCT image intensities of healthy tissue regions and

cancer in human axillary lymph nodes.................................................................. 55

Table 4-1 A summary of publications reporting OCT measurement of optical

properties in tissues and phantoms. ...................................................................... 70

Table 4-2 Attenuation coefficient ranges for tissue regions in healthy and reactive

excised human axillary lymph nodes. .................................................................... 93

Table B-1 Newport neutral density filter wheel 5214 optical density calibration. ........ 160

Table C-1 Compositon and properties of the microsphere suspension phantoms

used to select processing parameters and verify our attenuation

measurement procedure........................................................................................... 165

xv

Acknowledgements It would not have been possible for me to write this thesis without the help and support of others. I would like to

express my deepest thanks to the following people.

Firstly, to my supervisor David Sampson, who encouraged me to undertake a Ph.D at OBEL and who has

been a great mentor and guide to me throughout my candidature. I have so greatly appreciated his academic

guidance and unceasing support and encouragement along the way.

My sincere thanks to Robert McLaughlin at OBEL, who made it possible for me to pursue this exciting

project. I have learned much in the way of pragmatism, professionalism and generosity from working with him.

I could not have persevered through the many challenges that this project has presented without the

friendships and shared strivings of my fellow Ph.D’ers. A big thanks to Blake Klyen and Andrea Curatolo for the

many coffee breaks, fun times and invaluable discussions we have shared. Thank you to Yih Miin Liew, Kelsey

Kennedy and Shaghayegh Eshaghian for being the ladies I could look to for support, encouragement and

inspiration, and also to Lixin Chin, Xiao Jie Yang, and Peijun Gong for your shared enthusiasm for biomedical

engineering. To past Ph.D students: Tim Hillman, Tom Gutzler, Julian Armstrong and Steve Adie, thanks for

being so helpful and encouraging from the very beginning and for the many discussions that have guided me

along the way.

Thanks also to the OBEL postgrads and staff – it has been a privilege to work with such a successful and

internationally competitive team. I have very much enjoyed working alongside Dirk Lorenser and Bryden Quirk

who have been a tremendous source of knowledge and inspiration for me on the practical engineering aspects of

my project. Thank you to Liz Albert and especially to Tracey O’Keefe who has always been there to provide me

with the assistance and reassurance I needed to overcome any hurdle.

It has been a privilege to work with the brilliant clinicians and pathologists at Sir Charles Gairdner Hospital

and at Pathwest. My thanks to Christobel Saunders, Peter Robbins and Ben Wood for the many helpful

discussions on the clinical and histology aspects of the project and for their support throughout.

Thank you to Steve Jacques from the Oregon Health & Science University, who provided the invaluable

direction that expanded the focus of my project and who assisted me with some challenging theory along the way.

I could not have achieved this goal without the support of the William and Marlene Schrader Postgraduate

Scholarship for research in Biomedical Engineering at UWA, for which I am very grateful. I wish also to thank

the School of Electrical, Electronic and Computer Engineering, and the UWA Graduate Research School for

their financial support and travel funding.

Finally, to my wonderful family for their unfailing support during the toughest of times – we have been

through so much together over the past year especially. To my wonderful mum, Bianca, who has proven to be the

source of all good advice, and to my beautiful sisters Diana, Christine and Emily, and brothers-in-law Josh and

Adrian – no words can express how much I appreciate the support and strength each of you has given me. I hope

that I do justice to the memory of our wonderful Dad – it is a testament to his incredible wisdom, integrity and

diligence that we have each been able to achieve what we have.

Above all, my greatest thanks and praise goes to my Lord and Saviour Jesus Christ, who alone makes all

things possible and gives joy and strength in times of need. In you “are hid all the treasures of wisdom and

knowledge” (Colossians 2:3).

xvii

Statement of Contribution This thesis is an account of the research that I, Loretta Scolaro (LS), have undertaken during

my Ph.D. candidature at the School of Electrical, Electronic and Computer Engineering at

The University of Western Australia. Except as otherwise indicated here, the work described

is my own.

This thesis is primarily derived from five published (four full-length and one Letter), co-

authored journal articles1. Throughout this thesis, these publications are referred to by their

associated number in the bibliography on Page 169. Sections 3.3, 5.2 and 5.4 are reproduced

journal publications included without modification in content but the formatting has been

adapted to match the style of this thesis. Section 4.3 is an amalgamation of two journal

publications, which have been included with some modification of content. The details of

these publications and the locations where they appear are stated below. The contribution of

LS to each publication is given as a percent in brackets next to the author name and in the

descriptions below. LS is the sole author of the remaining sections of this thesis.

�• R. A. McLaughlin, L. Scolaro (30%), P. Robbins, S. Hamza, C. Saunders, and D. D.

Sampson, "Imaging of human lymph nodes using optical coherence tomography:

Potential for staging cancer," Cancer Res. vol. 70, no. 7, pp. 2579-2584, 2010

(Chapter 3: Section 3.3)

R. A. McLaughlin (RAM) is the principal author of the publication that forms Section 3.3.

Initiation of this research began with discussions between D. D. Sampson (DDS), LS, RAM

and C. M. Saunders (CMS). The contributions of LS to this publication were: to set-up and

optimise the OCT system used for imaging; to procure lymph nodes from pathologists at

PathWest Laboratory Medicine together with RAM; to assist in establishing a successful

protocol for performing OCT imaging of lymph nodes; to assist in interpreting the OCT

images obtained; and to scan histology slides using the digital slide scanner (Aperio

ScanScope) at The Centre for Microscopy, Characterisation and Analysis, Queen Elizabeth II

Medical Centre. CMS and S. Hamza were the surgeons responsible for performing the breast

cancer surgery at Sir Charles Gairdner Hospital. P. D. Robbins (PDR) was the pathologist

responsible for preparation and histopathological assessment of excised lymph nodes and for

providing expertise in clinical pathology. Ethics applications for research on human tissue

1 This thesis was prepared according to The University of Western Australia’s guidelines on “Thesis as a series of papers”

(http://www.postgraduate.uwa.edu.au/students/thesis/series. Accessed: 1 March, 2013)

xviii

were prepared by RAM1 who also wrote the visualisation software used to view and interpret

OCT images of lymph nodes in this section and subsequently in Sections 3.4, 4.3 and 5.2.

DDS was responsible for the conceptual development, design of the experimental programs,

overview of the analysis of results and drafting of the manuscript.

�• R. A. McLaughlin, L. Scolaro (20%), P. Robbins, C. Saunders, S. L. Jacques, and D. D.

Sampson, "Parametric imaging of cancer with optical coherence tomography," J. Biomed.

Opt. vol. 15, no. 4, art. 046029, 2010. (Selected for inclusion in the Virtual J. for

Biomedical Optics , vol. 20, no. 5, 2010).

(Chapter 4: Section 4.3.2 and part of Section 4.3.1) AND

�• L. Scolaro (60%), R. A. McLaughlin, B. R. Klyen, B. A. Wood, P. Robbins, C. M.

Saunders, S. L. Jacques, and D. D. Sampson, "Parametric imaging of the local attenuation

coefficient in human axillary lymph nodes assessed using optical coherence tomography,"

Biomed. Opt. Express, vol. 3, no. 2, pp. 366-379, 2012.

(Chapter 4: Sections 4.2.1, 4.3.3 and part of Section 4.3.1)

The work in both of these publications follows from that of the previous publication.

Therefore, many of the responsibilities and contributions are shared here. RAM is the

principal author of the first publication above from which half of Section 4.3.1 and all of

Section 4.3.2 were derived. The contributions of LS to this publication were the same as for

the previous publication, with the additional role of preparation and imaging of microsphere

tissue phantoms. S. L. Jacques (SLJ), initiated the project for optical property measurement

and provided the background theory for data analysis. Additional source code for data

processing and analysis was written by RAM.

LS was the principal author of the second publication above from which all of Section

4.2.1, half of Section 4.3.1 and all of Section 4.3.3 are derived. OCT images for these sections

were recorded, registered and interpreted by LS, who also conducted the additional

experimental procedures for calibration of the system and preparation and imaging of

microsphere phantoms for data correction and validation. Source code for data processing and

analysis was written by LS, with some early versions of the algorithms initially written by SLJ.

B. R. Klyen assisted in OCT system optimisation and system upgrade. B. A. Wood provided

additional expertise in pathology for histology assessment. DDS was responsible for

conceptual development and drafting of manuscripts as for the previous publication.

1 Ethics applications for all studies in this thesis were approved by both The University of Western Australia and Sir Charles Gairdner

Hospital Human Research Ethics Committee. Reference numbers: RA/4/1/5745 and 2007-152 respectively.

xix

�• R. A. McLaughlin, B. C. Quirk, A. Curatolo, R. W. Kirk, L. Scolaro (10%), D. Lorenser,

P. D. Robbins, B. A. Wood, C. M. Saunders, and D. D. Sampson, "Imaging of Breast

Cancer With Optical Coherence Tomography Needle Probes: Feasibility and Initial

Results," IEEE J. Sel. Topics Quantum Electron. vol. 18, no. 3, pp. 1184-1191, 2012.


The contribution of LS to the publication forming Section 5.2 was less significant, but this

publication has been included for completeness of the thesis argument and as an important

connection to the workflow in the next section. Initial OCT needle probe fabrication was

performed by LS in conjunction with B. C. Quirk (BCQ) who was also responsible for

motorising the rotation and pullback scanning mechanisms. A. Curatolo was responsible for

the system setup used for breast tumour imaging. D. Lorenser (DL) characterised the GRIN

fibre probes and developed the procedures and algorithms to do so. R.W. Kirk (RWK)

developed the software and interfaces for data acquisition and visualisation. The contribution

of the remaining authors was the same as for the previous publication.

�• L. Scolaro (50%), D. Lorenser, R. A. McLaughlin, B. C. Quirk, R. W. Kirk and D. D.

Sampson, "High-sensitivity anastigmatic imaging needle for optical coherence

tomography," Opt. Lett. vol. 37, no. 24, pp. 5247-5249, 2012. (Also selected for inclusion

in the Virtual J. for Biomedical Optics, vol. 8, no. 1, 2013).


LS is the principal author of the publication in Section 5.4. LS was responsible for fabrication

and characterisation of the needle probe, testing of the needle OCT system, and recording

needle OCT images of biological tissue. DL was responsible for needle design and for working

out initial steps of probe construction through discussions with engineers at Vytran (the

company supplying the filament fusion splicer used to make the probes). BCQ was

responsible for building the OCT system and ensuring its correct operation. RWK developed

the software for data acquisition and 3-D image reconstruction and visualisation. DL and

RWK developed the scripts for processing acquired data before image reconstruction. All data

acquisition, reconstruction, processing and interpretation was performed by LS.

Student’s signature: …………………………………………………………………………..Date:…………………...

Co-ordinating supervisor’s signature: ………………………………………………….Date:…………………...

xxi

List of publications The following is a list of publications arising from this thesis. PDFs and online multimedia

material of some of these publications are available on the website of the Optical+Biomedical

Engineering Lab (OBEL): http://obel.ee.uwa.edu.au/publications/. Fully refereed journal articles:

1. L. Scolaro, D. Lorenser, R. A. McLaughlin, B. C. Quirk, R. W. Kirk and D. D.

Sampson, "High-sensitivity anastigmatic imaging needle for optical coherence

tomography," Optics Letters vol. 37, no. 24, pp. 5247-5249, 2012. (Selected for

inclusion in the Virtual Journal for Biomedical Optics , vol. 8, no. 1, 2013).

2. L. Scolaro, R. A. McLaughlin, B. R. Klyen, B. A. Wood, P. Robbins, C. M. Saunders, S.

L. Jacques, and D. D. Sampson, "Parametric imaging of the local attenuation

coefficient in human axillary lymph nodes assessed using optical coherence

tomography," Biomedical Optics Express, vol. 3, no. 2, pp. 366-379, 2012.

3. R. A. McLaughlin, B. C. Quirk, A. Curatolo, R. W. Kirk, L. Scolaro, D. Lorenser, P. D.

Robbins, B. A. Wood, C. M. Saunders, and D. D. Sampson, "Imaging of Breast Cancer

With Optical Coherence Tomography Needle Probes: Feasibility and Initial Results,"

IEEE Journal of Selected Topics in Quantum Electronics vol. 18, no. 3, pp. 1184-1191,

2012.

4. R. A. McLaughlin, L. Scolaro, P. Robbins, C. Saunders, S. L. Jacques, and D. D.

Sampson, "Parametric imaging of cancer with optical coherence tomography," Journal

of Biomedical Optics vol. 15, no. 4, art. 046029, 2010. (Selected for inclusion in the

Virtual Journal for Biomedical Optics , vol. 20, no. 5, 2010).

5. R. A. McLaughlin, L. Scolaro, P. Robbins, S. Hamza, C. Saunders, and D. D. Sampson,

"Imaging of human lymph nodes using optical coherence tomography: Potential for

staging cancer," Cancer Research vol. 70, no. 7, pp. 2579-2584, 2010.

6. B. F. Kennedy, S. Loitsch, R. A. McLaughlin, L. Scolaro, P. Rigby, and D. D. Sampson,

"Fibrin phantom for use in optical coherence tomography," Journal of Biomedical

Optics vol. 15, no. 3, art. 030507, 2010 (This article was featured on the website

Biophotonics, September 2010, http://www.photonics.com/Article.aspx?AID=44334).

xxii

Selected conference papers:

Key: International | ^ Domestic | † Full paper| § Abstract| ‡ Refereed| ! Unrefereed or

abstract refereed.

7. ^§! L. Scolaro, R. A. McLaughlin, B. R. Klyen , B. A. Wood, P. D. Robbins, C. M.

Saunders and D. D. Sampson, "Parametric imaging of local attenuation coefficients in

human axillary lymph nodes using optical coherence tomography" presented at the

10th Asia-Pacific Microscopy Conference, Perth, WA, 5-9 February, 2012.

8. §! L. Scolaro, B. R. Klyen, R. A. McLaughlin, P. Robbins, C. M. Saunders, S. L.

Jacques, and D. D. Sampson, "Optical property measurement with 3D-OCT to

differentiate soft tissues," presented at SPIE Photonics West BiOS, Biomedical

Applications of Light Scattering V, San Francisco, USA, 22-24 January, 2011.

9. †! R. A. McLaughlin, L. Scolaro, P. Robbins, C. Saunders, S. L. Jacques, D. D.

Sampson, “Tissue differentiation in human lymph nodes using parameterized optical

coherence tomography”, presented at SPIE Photonics West BiOS, Coherence Domain

Optical Methods and Optical Coherence Tomography in Biomedicine XIV, San

Francisco, USA, 23-30 January, 2010.

10. ! B. C. Quirk, R. A. McLaughlin, L. Scolaro, P. Robbins, C. Saunders, D. D. Sampson,

“3D-OCT imaging of ex vivo human tissue using a novel rotating needle probe”,

presented at SPIE Photonics West BiOS, Advanced Biomedical and Clinical

Diagnostic Systems VIII, San Francisco, USA, January 23-30 2010.

11. †‡ R. A. McLaughlin, L. Scolaro, P. Robbins, C. Saunders, S. L. Jacques, and D. D.

Sampson, "Mapping tissue optical attenuation to identify cancer using optical

coherence tomography," in Medical Image Computing and Computer-Assisted

Intervention – MICCAI 2009. (Eds. G. Z. Yang, D. Hawkes, D. Rueckert, A. Noble,

and C. Taylor, Springer Berlin), Lect. Notes Comput. Sci., 5762, pp. 657-664, 2009.

12. §! L. Scolaro, T. Wienhold, R. A. McLaughlin, B. R. Klyen, S. L. Jacques, and D. D.

Sampson, "Measuring tissue optical properties with needle probes using optical

coherence tomography," presented at SPIE Photonics West BiOS, Biomedical

Applications of Light Scattering III, San Jose, USA, 24-29 January, 2009.

13. †! R. A. McLaughlin, L. Scolaro, B. R. Klyen, S. Hamza, P. Robbins, C. Saunders, D.

D. Sampson, “Optical coherence tomography of human breast lymph nodes:

microstructures and metastasis”, presented at SPIE Photonics West BiOS, Coherence

xxiii

Domain Optical Methods and Optical Coherence Tomography in Biomedicine XIII,

San Jose, USA, 24-29 January 2009.

14. §! R. A. McLaughlin, L. Scolaro, B. R. Klyen, S. Hamza, P. Robbins, C. Saunders, D.

D. Sampson, “Lymph node micro-architecture can be imaged using Optical Coherence

Tomography”, presented at the 31st Annual San Antonio Breast Cancer Symposium,

San Antonio, USA, 10-14 Dec 2008, and published in Cancer Research, vol. 69, no. 2,

pp. 102S-102S, Supplement, Jan 15, 2009.

15. †! R. A. McLaughlin, L. Scolaro, B. R. Klyen, J. J. Armstrong, S. Hamza, P. Robbins,

C. Saunders, D. D. Sampson, “Can normal lymph node architecture be characterised

by optical coherence tomography?”, Proc. SPIE vol. 7139: presented at the 1st

Canterbury Workshop on Optical Coherence Tomography and Adaptive Optics,

Canterbury, UK, 8-10 September 2008, paper 19.

Other publications: 16. R. A. McLaughlin, B. C. Quirk, D. Lorenser, X. Yang, B. Y. Yeo, A. Curatolo, K. M.

Kennedy, L. Scolaro, R. W. Kirk, and D. D. Sampson, “A microscope in a needle”,

Optics and Photonics News, vol. 23, no. 12, p. 40, December 2012. (This article was

selected for the 2012 special issue “Optics in 2012”. It also featured in a video highlight

of the issue, http://www.osa-opn.org/home/articles/volume_23/december_ 2012).

xxv

List of Acronyms ALND Axillary lymph node dissection

CNB Core-needle biopsy

CT X-ray computed tomography

DAQ Data acquisition card

FDOCT Fourier-domain OCT

FFT Fast Fourier transform

FNAC Fine needle aspiration cytology

FWHM Full width at half-maximum

GRIN Graded- or gradient-index

H&E Haematoxylin and eosin

ITC Isolated tumour cells

MEMS Micro-electro-mechanical systems

MRI Magnetic resonance imaging

NA Numerical aperture

NCF No-core fibre

NDF Neutral density filter

Near-IR Near-Infrared

OCT Optical coherence tomography

OD Optical density

PBS Phosphate-buffered saline

PET Positron emission tomography

SDOCT Spectral-domain OCT

SLNB Sentinel lymph node biopsy

SMF Single-mode fibre

SNR Signal-to-noise ratio

SSOCT Swept-source OCT

TDOCT Time-domain OCT

TIR Total internal reflection

US Ultrasound

VOA Variable optical attenuator

WD Working distance

1

Chapter 1

Introduction 1

One pathway for the metastasis (or spread) of breast cancer to other parts of the body is

through the lymphatic system [1]. The lymphatic system transports tissue fluid, called lymph.

Its primary function is to provide an accessory route for lymph to be returned from tissues

throughout the body to the blood stream. Lymphatic organs play an important part in the

immune system: lymph nodes are the organs that facilitate the body’s immune response by

filtering the lymph as it circulates. It follows that cells of metastatic cancer usually appear first

in the lymph nodes, specifically in the lymph nodes that are closest to the site of a primary

tumour. For breast cancer, the lymph nodes most commonly involved in metastasis are those

located in the axilla (or underarm).

The presence of metastatic cancer in axillary lymph nodes is an important prognostic

indicator for patients with breast cancer [2]. Both the number of nodes involved and the

extent to which nodes are affected are important factors to determine. Nodes that contain

metastatic cancer (referred to as malignant or involved nodes) indicate the presence of a

malignant tumour and prognosis for patients with axillary metastasis is much worse than for

those without metastasis. In evaluating lymph node involvement, it has become standard

practice that only the sentinel (or first) lymph node on the lymphatic pathway that drains the

tumour is assessed. This procedure is called sentinel lymph node biopsy (SLNB) [3]. If the

sentinel node is not involved, prognosis is good and the remaining nodes are not removed. If it

is determined that one or more axillary lymph nodes are involved (contain metastatic cancer),

it is usually recommended that at least two thirds of all axillary nodes are removed as part of

the staging and treatment process [3, 4]. This procedure, called axillary lymph node dissection,

is generally performed at the same time that the primary tumour is excised.

Currently, the accepted gold-standard method to determine lymph node involvement is

excision and postoperative diagnosis via histopathology. However, removal of lymphoid tissue

often results in negative side effects, the most severe of which is swelling of the arm, called

lymphoedema [5]. As excision is currently required to make a diagnosis, patients may suffer

side effects unnecessarily if the excised nodes are later found to be negative for the presence of

2 Chapter 1 Introduction

cancer. In addition, since biopsy is usually performed while the patient is undergoing primary

tumour excision, the time available for making a diagnosis is very limited.

Clearly, a reliable alternative method of biopsy that does not require lymph node removal

would be preferable to the current standard method of excisional biopsy. Such an in situ

biopsy method could improve lymph node diagnosis in several ways, in particular:

�• in situ biopsy could be used to identify healthy (negative) sentinel lymph nodes during

breast cancer surgery, thereby eliminating the need for their removal. This could reduce

the number of healthy lymph nodes excised and avoid unnecessary risk of lymphoedema

or other side effects;

�• a pre-operative implementation of in situ biopsy could establish the presence of positive

axillary nodes prior to surgery. This could improve therapeutic and surgical planning for

patients, and enable more informed decisions to be made in relation to neo-adjuvant

systemic therapy (i.e., chemotherapy); and

�• for more definitive pre-operative diagnosis, in situ assessment could be employed

alongside traditional pre-operative excisional needle biopsy techniques (such as core

needle biopsy), in order to guide accurate sampling of nodal tissue and reduce the false

negative rates of these techniques.

Some alternative clinical imaging methods for in situ lymph node assessment have already

been investigated, including: mammography, X-ray computed tomography, positron emission

tomography, magnetic resonance imaging and ultrasonography [6]. However, these methods

have not yet proven to be sufficiently accurate or suitable for routine use. Optical diagnostic

techniques, on the other hand, may have greater potential for detecting the signature of

cancer, due to their high resolution, high detection sensitivity, low toxicity and potential for

real time assessment.

Light has long been recognised as an important tool in medicine. One of the earliest Nobel

prizes was awarded to Niels Ryberg Finsen in 1903 for his use of light therapy for the

treatment of skin disease. In modern medicine, the use of light is still important for

therapeutic applications in dermatology; however, light is more commonly used in medicine

for diagnosis. For example, optical microscopy is the fundamental tool for diagnosis of disease

in excised tissue, as it enables histological evaluation of tissue pathology. For in situ optical

assessment of tissue, the term “optical biopsy” has come into common use. “Optical biopsy”

refers to an optical measurement tool (most commonly, but not limited to, a type of

spectroscopy) that is used to diagnose tissue pathology in real-time. Although the term


‘biopsy’ actually refers to the removal of tissue, optical biopsy is commonly understood to

represent biopsy that does not require tissue removal. For example, diffuse optical tomography

and optical spectroscopy have been shown to reveal pathological changes in the structure and

function of breast tissue in vivo [7, 8]. The properties of light – its intensity, wavelength,

phase/group delay, degree of coherence (spatial and temporal) and polarisation – are

fundamentally affected by the micro- and macro-composition of the tissue through which it

propagates. Understanding and detecting the changes in these properties is the aim of optical

biopsy. Although in situ optical methods of tissue diagnosis are relatively new, and have not

yet been implemented clinically, they are gaining significant attention in medical research.

Optical coherence tomography (OCT) is one such modality that is being explored as a

tool for in situ optical biopsy [9]. First developed in the early 1990’s, OCT is a high-

resolution, three-dimensional optical imaging modality that measures the intensity of light

elastically backscattered from a tissue, analogous to the way ultrasound imaging measures the

reflection of sound [9, 10]. As the speed of light is much faster than that of sound, OCT

requires a more complex method of detection. Therefore, unlike ultrasound imaging, OCT

uses interferometry and a low-temporal-coherence light source to pinpoint the depth-location

of reflected light. It does so through the generation of a coherence gate, which strongly

discriminates detection of singly scattered from diffusely scattered light. The main advantage

of imaging with light is that it enables a higher resolution than imaging with sound: on the

scale of 10 – 20 μm for OCT compared to typically >100 μm for ultrasound imaging.

The resolution of OCT fills the gap between traditional microscopic techniques (such as

confocal microscopy) and clinical imaging techniques (such as magnetic resonance imaging,

computed tomography, and ultrasound) (see Figure 2-7 in Section 2.3, ‘Resolution and

imaging depth of OCT compared with other imaging modalities’). OCT resolution is lower

than that of confocal microscopy, but this is compensated for by its increased penetration

depth, which can be up to ten times greater, at around 2 to 3 mm compared to 0.2 mm. This is

possible due to the high detection sensitivity of OCT as a result of the amplification obtained

when multiplying the weak sample signal with a strong reference signal. In addition, the axial

resolution of OCT is determined by the coherence gate and is therefore decoupled from the

lateral resolution. This permits the use of a low NA objective making axial translation of the

objective beam within the sample unnecessary. Although the penetration depth of OCT does

not reach that of clinical imaging modalities (such as MRI or ultrasound), its non-invasive,

high-resolution, three-dimensional imaging capability, as well as its low cost and high

portability, make it a very attractive tool for clinical tissue assessment.


It is also possible to achieve ultra-high resolution with OCT, approaching 1 μm both

axially and laterally [11]. Unlike traditional microscopy, axial resolution with OCT is

decoupled from the properties of the imaging beam. This means an increase in axial resolution

does not come at the expense of a reduction in depth of field and, consequently, penetration

depth (see also Section 2.3.1, ‘Basic principles of OCT’).

One application where the advantages of OCT have been particularly beneficial is

ophthalmology. OCT is now commonly accepted as a clinical standard for visualisation of

structure and disease in the retina and for imaging of the cornea and anterior segments of the

eye [12, 13]. A literature search conducted in Elsevier’s Scopus database1 provides a useful

snapshot of the major research areas for OCT applications in biology and medicine over the

past ten years. A summary of the search results is set out in the chart in Figure 1-1 below. The

search was conducted for publications within the past ten years, with a title containing the

words “optical coherence tomography” and with words relating to specific OCT biomedical

applications, as indicated in the chart. It is evident from the graph that ophthalmology has

consistently been the most researched OCT biomedical application, with more than 1776

research publications.

Figure 1-1. Bar chart of a snapshot of results from a Scopus search of selected OCT biological applications over the past ten years. Searches were conducted on the publication title and only included English results. Search terms used were as follows: “optical coherence tomography” AND (terms in brackets under each heading) e.g. "optical coherence tomography" AND (breast OR mamm*), where the use of a * indicates a search for all words containing the truncated term.

1 http://www.scopus.com


In easy-to-access superficial tissues such as skin, the use of OCT has been explored for

dermatological assessment of skin cancer, scar assessment and other diseases [14, 15]. As the

resolution, penetration depth and imaging speed of OCT has improved over the past ten

years, it has steadily gained increasing attention in this field. Similarly, in dentistry, the

benefits of OCT are fast being realised, particularly because enamel is highly transparent at

near-infrared (near-IR) wavelengths [16].

OCT imaging of less accessible surfaces within the hollow organs of the body is possible

using endoscopic fibre-optic probes. In cardiology in particular, endoscopic OCT is finding a

niche for imaging atherosclerotic plaques of coronary arteries [17]. Recent growth in

commercially available intravascular OCT imaging systems has meant that cardiology has

been the second most published clinical application of OCT, as shown in Figure 1-1 above.

Recent advances in fibre optics and photonics have improved the speed and quality of

imaging and have expanded the range of applications for endoscopic OCT. Applications

where endoscopic OCT is emerging as a useful clinical imaging tool include in imaging of the

gastrointestinal tract (gastroenterology) and respiratory tract (pulmonary medicine) [18, 19].

For tissues of the gastrointestinal tract (including the oesophagus, colon and bile/pancreatic

duct), endoscopic OCT is especially useful since it enables visualisation of subsurface tissue

morphology. This is a significant advantage over conventional endoscopy, which only enables

imaging of surface features. In pulmonary medicine, endoscopic OCT has been useful for

visualising abnormal epithelia in the proximal airway and for imaging the size and shape of the

upper airway. Other applications for endoscopic OCT have included urology (bladder and

prostate), gynaecology and laryngology.

Endoscopic OCT is not limited to superficial tissues of hollow organs. The recent

development of highly miniaturised endoscopes using small fibre-optic components has

enabled OCT imaging to be translated into needles. In situ OCT imaging of solid tissues with

needle probes has shown much promise, although it has not yet been demonstrated in a

clinical study. The application of OCT needle probes has only recently begun to be explored

for imaging of tissues in the breast [20], lung [21], brain and pancreas. Imaging of tissues

within the brain for applications in neurology in animal models have particularly made

progress over the past two years [22].

One of the greatest challenges for extending clinical applications of OCT is to find

improved contrast mechanisms that can better differentiate tissue pathology. This challenge is

being addressed by the development of functional extensions of OCT, which can provide

additional information not contained in structural images. Doppler OCT is one such


functional extension that can detect particle motion in order to image tissue blood flow. This

enables visualisation of vasculature and blood flow velocity, which can be superimposed on

structural OCT images. Since many pathologies exhibit an abnormal vasculature, Doppler

OCT may prove useful in a number of applications including ophthalmology, dermatology,

gastroenterology and neurology [23]. Polarisation sensitive OCT (PS-OCT) is another

extension of OCT that measures the polarisation state of reflected light to detect changes in

birefringence and dichroism [24]. PS-OCT has been investigated for imaging in birefringent

tissues such as skin and muscle. Spectroscopic OCT combines spectroscopic analysis with

OCT to obtain depth-resolved tissue absorption spectra; such spectra can be used to measure

localised tissue oxygen saturation [25]. More recently, OCT elastography has been

investigated to exploit the variations in elasticity of healthy tissue and cancer in the breast

[26]. Other means of contrast, including the use of molecular probes, spectral properties and

optical properties, may also lend further capacity to OCT imaging for some applications [27].

1.1 Thesis motivation

The potential for OCT to perform optical biopsy is, therefore, well recognised in many

biological and medical applications. The use of OCT imaging for optical biopsy of cancer is a

promising, albeit challenging, area of research and there is an evident lack of exploration of

OCT for imaging of malignant lymph nodes. This thesis explores the potential of OCT in the

optical biopsy of malignant axillary lymph nodes.

A baseline study of OCT imaging of lymph nodes will be presented, along with two novel

developments that advance the capability of OCT imaging towards in situ optical diagnosis of

lymph node status. The first development uses an alternative means of OCT contrast for

differentiating healthy and malignant tissue, based on their differing optical properties. The

second explores the translation of OCT imaging hardware into miniature optical probes

inserted through hollow needles.

1.2 Structure of this thesis

Chapter 2 - Background

Chapter 2 provides the background on the role of lymph nodes in the spread of breast cancer.

The standard method for assessing the involvement of lymph nodes based on excision and

histological evaluation is described, as well as the possible side effects of lymphoid tissue


excision. A summary of alternative techniques is then presented, followed by an introduction

to OCT. The specifications of the OCT systems used in the investigations of this thesis are

also outlined.

Chapter 3 – OCT imaging of axillary lymph nodes

This chapter investigates the potential use of OCT (in its standard form) for imaging of

excised human axillary lymph nodes. To this end, the tissue composition of healthy and

malignant axillary lymph nodes is discussed, with reference to representative histology images

of healthy and malignant lymph nodes. A short summary of the current literature on OCT

imaging of lymph nodes is also set out. The following section, reproduced from [28], presents

an investigation into the ability of OCT to image lymph node micro-architecture and identify

metastasis, based on comparison with corresponding histology images. The chapter concludes

with a discussion of some of the shortcomings of intensity-based image contrast for

differentiating malignant and healthy lymph nodes, and the need for improved contrast.

Chapter 4 – OCT for optical property measurement in lymph nodes

In Chapter 4, we introduce and demonstrate a method for improving the contrast of OCT

imaging, based on extracting a tissue optical property, specifically the attenuation coefficient.

We visualise extracted attenuation coefficients in the form of en face parametric maps

(parametric images). There are various methods for modelling and analysing OCT image

intensity to extract the attenuation coefficient, and a literature review of current methods is

discussed. Our chosen method is subsequently described, based on the single-scattering model

of OCT image intensity. The next part of this chapter presents an amalgamation of two

previously published studies ([29] and [30]), which report on the use of the single-scattering

model to extract the relative and quantitative attenuation coefficients of healthy and

abnormal (including malignant) lymph nodes. These results support the use of parametric

imaging to aid interpretation of the OCT appearance of lymph nodes and facilitate their

diagnosis.

Chapter 5 – Needle-based OCT

The fifth chapter presents a development of the OCT imaging technology to achieve imaging

of tissue in situ, thereby enabling tissue analysis to be conducted without the need for excision.

This chapter describes the design and implementation of two novel OCT needle probes – the

second being an advanced version of the first. The material for these sections is reproduced

from our previous publications: [31] and [32]. The application of each needle is demonstrated


for imaging of excised lymph nodes and breast tumour tissue. Finally, we present the capability

of the second needle to measure optical properties of tissue. This work provides a proof of

principle demonstration that needle OCT parametric imaging can be used to differentiate

healthy and malignant lymph nodes.

Chapter 6 - Conclusion

The final chapter summarises our main findings, which are as follows: firstly, that OCT can

identify the micro-architectural appearance of healthy lymph nodes and that cancer has an

identifiable appearance in OCT images characterised by high scattering intensity; secondly,

that parametric imaging of the local optical attenuation coefficient, extracted from OCT

image data, can improve contrast and interpretation of OCT images and guide tissue

segmentation; and finally, that OCT parametric imaging can successfully be implemented in

optical needles and be used to differentiate human axillary lymph nodes and breast tumour

tissue. Based on these conclusions, we discuss the potential of OCT to achieve in situ lymph

node biopsy and discuss some alternative applications for this work that may ultimately prove

useful. We finish with a discussion of future work that would be required to realise our aim of

enabling practical and reliable in situ optical biopsy of axillary lymph nodes.

9

Chapter 2

Background 2

In this chapter we give the background of lymph node involvement in breast cancer and

explain the motivation for using optical coherence tomography (OCT) to assess the status of

axillary lymph nodes. We first discuss the importance of lymph node assessment for cancer

staging and the current method to diagnose cancer in excised lymph nodes by excision and

histopathological evaluation. We outline the side effects of lymph node removal including

lymphoedema and give a brief review of the medical imaging literature focusing on alternative

methods for lymph node assessment. Optical imaging by OCT is discussed as an alternative to

achieve lymph node ‘optical biopsy’. We discuss the advantages and disadvantages of OCT as

an optical biopsy tool and present the principles of OCT and a summary of the OCT systems

used in the investigations reported in this thesis.

2.1 Breast cancer and lymph node metastasis

Breast cancer is the most common form of cancer in women, accounting for almost one in

four female cancer diagnoses worldwide in 2008 [33]. Most breast cancers have their origins at

the terminal ductal lobular unit of the breast with cellular proliferation occurring within these

units and within the adjoining ductal structures. For early stage cancer, the tumour is usually

confined to these ducts and is referred to as ductal carcinoma in situ. Breast cancers progress

from in situ to invasive when malignant cells infiltrate the duct walls and invade the

surrounding breast tissue. Invasive breast cancer becomes metastatic when malignant cells

break away from the main tumour and spread to other organs of the body through either the

blood stream or the lymphatic system. The risk of malignant cells metastasising through the

lymphatic system is directly proportional to the size of the primary breast tumour [2]. Invasive

breast cancer is the second leading cause of cancer-related female deaths in Australia and the

US [34]. Accurate determination of the extent and spread of disease (or staging) is a critical

component of the care and management of individuals with breast cancer.

10 Chapter 2 Background

2.1.1 Disease progression

The lymphatic system provides one pathway for the dissemination of malignant cells in the

body, and lymph nodes represent a common site of metastatic involvement. Like the

circulatory system, the lymphatic system is made up of vessels that transport fluid. In the case

of the lymphatic system this fluid is lymph or tissue fluid that drains from nearby tissues.

Deposits of metastatic cells are, therefore, usually observed first in the regional lymph

nodes. Lymph nodes are a major component of the lymphatic system that lie in clusters or

chains along the lymphatic vessels at strategic locations where they drain nearby anatomic

regions, such as the neck, abdomen and axilla. They facilitate the body’s immune response

through mechanical filtration of the foreign bodies in the lymph and through recognition and

processing of antigens. The regional lymph nodes of the breast are those clustered in the axilla

as illustrated in Figure 2-1. There are about 20-35 lymph nodes in the region of the axilla and

this is usually the first anatomic site that is involved in breast cancer metastasis.

Figure 2-1. Illustration of breast cancer and its metastasis to lymph nodes. (a) Metastatic breast tumour spreads to the loco-regional lymph nodes in the axilla. The sentinel node is the first along the lymphatic pathway draining the tumour and is usually the first site where metastatic cancer is found. (b) Cancer metastasises to other organs of the body through the blood stream or through the lymphatic system. From: The website of the National Cancer Institute, “Breast Cancer Treatment”, 2012. [Online]. Available: http://www.cancer.gov/cancertopics/pdq/treatment/breast/Patient. [Accessed: 11 December 2012]. Used with permission from Terese Winslow, 2012.

2.1 Breast cancer and lymph node metastasis 11

2.1.2 Staging and treatment

Involvement of axillary lymph nodes in cancer metastasis provides the single most important

indicator for prognosis in early breast cancer [35]. Involved lymph nodes represent an invasive

tumour biology and, therefore, an increased risk of secondary cancer and mortality. The stage

of breast disease is classified according to tumour size and the degree of lymph node

involvement, with both the number and location of histologically positive lymph nodes

having prognostic significance. For staging of breast cancer: Stage 0 represents no nodal

involvement, Stages I and II may have some nodal involvement and Stages IIIA, IIIC and IV

always have nodal involvement [36]. The stage is indicative of the severity of disease with later

stages being more severe.

Where lymph nodes are involved, the standard recommendation for therapy is often

complete axillary clearance; that is, removal of all accessible axillary lymph nodes in order to

eliminate the nodes that may contain cancer and prevent further cancer spread.

2.1.3 Morphology of lymph nodes

Lymph nodes are small and roughly bean-shaped organs, varying in size from 2 mm to 20 mm.

They lie within a supporting framework of adipose tissue that surrounds the node and usually

occupies the concave space at the central hilum where blood and lymphatic vessels enter and

exit. A diagram of the typical morphology of a lymph node is shown in Figure 2-2.

Lymph nodes have a complex architecture in which a variety of cell populations are

arranged in distinct interfacing compartments. The entire node is enclosed by a fibrous

capsule of dense connective tissue with trabeculae that extend deep into the parenchyma as

part of the supporting stromal framework. The lymph node parenchyma consists mainly of

small immune cells called lymphocytes in various stages of maturation, as well as supporting

cells, reticular fibres (made up of fine collagen fibrils) and blood vessels. Lymph nodes can be

roughly divided into the following compartmental regions or zones as indicated in Figure 2-2:

cortex (light blue in Figure 2-2), in which primary and secondary follicles are generally

located; paracortex (light green), situated just below the cortex and between cortical follicles;

and medulla, which lies deeper in the node and consists of medullary cords (dark green) and

medullary sinuses (pale yellow). The medullary sinuses eventually join to form the efferent

lymphatic vessels.


Figure 2-2. Basic structure of a typical lymph node. The lymph node is roughly bean shaped with an indentation at the hilum where blood and lymphatic vessels exit. It is cushioned by surrounding adipose tissue (not shown). The internal structures include a capsule of fibrous connective tissue (brown) and a parenchyma of lymphocytes and supporting cells compartmentalised into cortex (light purple) with cortical follicles (dark purple), paracortex(light green) and medullary cords (dark green) and sinuses. Adapted from: Stevens A, Lowe J. Histology. London:Gower Medical Publications, 1992, pp. 90-95.

Lymphatic flow into the lymph node occurs through afferent lymphatic channels that

enter at the capsule and drain through subcapsular sinuses into the medullary sinuses.

Outward flow of lymph occurs through the efferent lymphatic channels at the hilum. It is

often the case, therefore, that early metastatic deposits are identified first in the subcapsular

sinuses [37]. However, metastatic cells can also spread through the whole node to completely

replace normal lymph node tissue. Malignant nodes containing metastatic tumour are

generally larger and harder than healthy benign reactive lymph nodes. More detail on the

cellular morphology of healthy and malignant lymph nodes is given in Chapter 3, in Section

3.2.2 and Section 3.3.3, respectively.

2.1.4 Standard histopathological evaluation

The current gold standard method for identification of metastasis in lymph nodes is

histopathological evaluation. To facilitate accurate assessment, excised lymph nodes are first

bisected into 2 mm portions along the long axis. Each portion is then evaluated by

histopathology. Standard histological preparation involves fixation, dehydration, mounting

and thin slice sectioning of the tissue, followed by appropriate staining to visualise the

morphology of cells at high resolution using an optical microscope. The most widely used

stain in pathological diagnosis, and the one utilised for the work in this thesis, is haematoxylin

and eosin (H&E) that stains basophilic structures (such as nuclei) blue, and eosinophilic

structures (such as cell cytoplasm and collagen) pink.

2.1 Breast cancer and lymph node metastasis 13

2.1.5 Side-effects of histopathological evaluation

The standard method for lymph node diagnosis is histopathological evaluation. However,

histopathology requires nodes to be excised before they can be assessed. In addition to causing

trauma for the patient, there are significant side effects associated with lymph node removal,

including chronic lymphoedema, restricted arm movement, partial sensory loss (due to

sensory nerve injury) and pain. Lymphoedema is a swelling of the arm, breast or chest wall that

is caused by an accumulation of lymph in the tissue. An example of the effect of lymphoedema

on a patient after lymph node removal is shown in Figure 2-3. Medical management of the

condition can include intermittent pneumatic compression, decongestive physiotherapy and

elevation. These options usually require lifetime behaviour modification. In a survey of the

literature, Petrek et al. found that patients undergoing axillary lymph node dissection

(ALND) developed lymphoedema in 6% to 36% of cases, depending on such factors as weight

gain, node status and arm infection/injury [5].

Figure 2-3. Severe lymphoedema in the right arm of a patient after ALND for breast cancer staging. From: Women's Health & Education Centre, “Breast Cancer Surgical Treatment Complications & Lymphoedema”, 2011. © WHEC. [Online]. Available: http://www.womenshealthsection.com/content/gyno/gyno005.php3. [Accessed 11 February 2013].

2.1.6 Sentinel lymph node biopsy

Historically, nodal involvement has been determined by ALND, in which at least two thirds

of lymph nodes in the axilla are removed and evaluated by histopathology. Although ALND

remains the "gold standard" for sensitivity and accuracy of detection, it carries a higher

morbidity than the alternative and now widely accepted technique of sentinel lymph node

biopsy (SLNB) [38, 39]. SLNB requires the removal of only the sentinel lymph node, which is

the first draining lymph node on the direct lymphatic pathway from a primary tumour,

although there may be more than one sentinel node. If the sentinel node does not contain

metastasis, ALND can be avoided with minimal increased risk of mortality [39]. Thus, SLNB

is an alternative to ALND that significantly reduces morbidity associated with lymphoedema.


The location of the sentinel lymph node is determined by ‘mapping’ through the injection

of a radioactive colloid and/or a blue dye (patent blue) near the site of the tumour [3]. The

radioactive colloid is injected several hours prior to surgery and a lymphoscintigram image (an

example of which is shown in Figure 2-4) roughly locates the sentinel lymph node. During the

surgery, a hand-held gamma probe is used to accurately locate the ‘hot’ node. Patent blue dye

is a complementary method that is used to visually identify the node by its blue appearance

after injecting the dye near the tumour at the commencement of surgery.

Figure 2-4: Lymphoscintigram transmission images taken prior to surgery. Anterior (A) and right lateral (B) images taken 30 minutes after injection of the radionuclide dye in the left breast (arrow). Focal uptake can be observed in the sentinel lymph node (dashed arrow). Originally published in JNM: Pandit-Taskar, N. et al., “Organ and fetal absorbed dose estimates from 99mTc-sulfur colloid lymphoscintigraphy and sentinel node localisation in breast cancer patients”, J. Nucl. Med., 2006, vol. 47:1202-1208, by the society of Nuclear Medicine and Molecular Imaging Inc.

Diagnosis of sentinel node status is usually performed while the patient is undergoing

surgery, so that the surgeon can proceed with a full therapeutic axillary clearance if the node is

found to be positive for metastasis. Therefore, to minimise surgery time and risk to the

patient, fast assessment of the node is required and this is usually accomplished using frozen

section histology or imprint cytology. These techniques typically require 15-30 minutes to

complete. For frozen section histology, a section is made by snap freezing the tissue to enable

slicing, followed by staining and viewing of the section under a microscope. Alternatively,

imprint cytology takes a cell ‘smear’, rather than a section, from an internal cut surface of the

node that is then microscopically examined. Although these techniques are fast to perform,

their sensitivity and specificity (53 – 75% and 98% – 100%, respectively [40, 41]) are much

lower than for standard histopathological assessment because of the error involved in bisecting

the node at the location of a metastatic deposit. In particular, micro-metastases, which are

defined as metastases less than 0.2 mm in size are often missed. Detection of micro-metastases

is a significant challenge for lymph node diagnosis, but there is also considerable debate as to

whether they are significant for prognosis [42]. In accordance with low sensitivity, there is a

recognised false negative rate of diagnosis with these techniques: in a survey of the literature

2.2 Alternative techniques for in situ assessment of lymph node status 15

Kim et al. found that an average of 7.3% of breast cancer patients with positive axillary nodes

were falsely diagnosed as negative at biopsy [43]. Thus, excised sentinel nodes are usually also

submitted for detailed histopathological analysis after surgery. Additional therapeutic ALND

(that is axillary clearance) may then be required in a second surgery based on the results of

histopathology; however, this occupies additional valuable surgery time and can cause

increased trauma for the patient.

The risk of lymphoedema after SLNB is lower than for ALND; however, there is still an

associated risk. In one study the occurrence of lymphoedema for SLNB patients was around

6.9% [44]. In the same study, 3.8% of patients who underwent SLNB reported restricted

movement and 8.6% reported feeling numbness at 6 months after surgery. Furthermore, only

42% of sentinel nodes contain metastases [43], indicating that uninvolved healthy lymph

nodes are excised in a significant proportion of cases.

2.1.7 Motivation for a new assessment technique

The motivation for an in situ intra-operative assessment technique is to avoid the unnecessary

excision of lymph nodes when the axilla is free of malignancy. A high accuracy in situ method

for assessment of node status prior to excision could reduce the rate of removal of negative

sentinel or axillary nodes and improve patient outcome and well being. This method could be

employed intra-operatively or pre-operatively. The advantage of an intra-operative assessment

technique is that it would work well with current SLNB procedures and, if fast enough could

also reduce the time required for lymph node assessment during surgery.

2.2 Alternative techniques for in situ assessment of lymph node status

SLNB is now commonly employed for breast lymph node assessment worldwide. However,

the search for a less invasive clinical method for axillary lymph node assessment is of increasing

interest and has led to research growth in a number of current and new modalities for lymph

node assessment. Alternative modalities for lymph node assessment may be broadly classified

into pre-operative needle biopsy techniques, standard clinical imaging techniques and novel

optical imaging techniques. The success of these modalities to assess the status of axillary

lymph nodes depends on their ability to identify either the molecular, morphological or

vasculature signature of metastasis with minimal side effects. Thus, we discuss the most


significant modalities here, in light of their diagnostic accuracy (sensitivity and specificity) for

detecting lymph node involvement.

2.2.1 Pre-operative needle biopsy

Pre-operative needle biopsy uses a hollow-core needle to obtain a sample of cells for fine-

needle aspiration cytology (FNAC) or a section of tissue for core-needle biopsy (CNB) [45-

48]. Extracted samples are examined by histopathological evaluation, similar to SLNB. For

FNAC, an aspirate of cells is taken from the node using a small diameter (21 – 23-gauge, 0.8 –

0.65 mm outer diameter) needle. The aspirate is smeared onto a microscope slide and the cell

morphology is assessed for cancer. Alternatively, CNB uses a larger needle (14-gauge, 2.1 mm

outer diameter) with an outer cutting cannula that is used to excise a small section of tissue

after insertion. The specificity for predicting positive nodes by either FNAC or CNB is very

high, close to 100%, since cancer cells are easy to identify under histopathology. On the other

hand, it is difficult to locate the nodes within the body using these techniques and there is

significant error in sampling diagnostically significant tissue, that is, tissue from the node as

opposed to external adipose or connective tissue. Thus, sensitivity can be very low ranging

from 57% to 87% for FNAC [45, 49] and from 40% to 94% for CNB [50, 51]. For CNB,

there is also the added complication that the procedure is invasive and problems can arise

from damaging major vessels and nerves [51]. Some work has been done to improve sampling

through better guidance. Both of these techniques must be performed under image guidance

using an external imaging modality to direct the needle to the location of the node. This is

usually accomplished by ultrasound imaging. Recent studies have looked into better mapping

of lymph nodes in order to improve guidance either using colour-Doppler ultrasonography or

photo-acoustic tomography with contrast agents [50].

2.2.2 Pre-operative imaging

Pre-operative imaging techniques may provide significant benefits since these have the

potential for non-invasive assessment of the axilla. The most notable current in vivo imaging

techniques being investigated for this purpose are: mammography, X-ray computed

tomography (CT), positron emission tomography (PET), ultrasound (US) and magnetic

resonance imaging (MRI) [6].

Mammography is the standard screening tool for initial detection of tumours of the breast.

A positive diagnosis is based on increased absorption of X-rays from high-density tumour

tissue and the presence of micro-calcifications. Mammographic imaging extended to the axilla

2.2 Alternative techniques for in situ assessment of lymph node status 17

detects pathological lymph nodes based rather on their size. However, contrast based on size is

not considered an accurate means of differentiating healthy and malignant lymph nodes and,

thus, the accuracy of mammography is very low [6, 52].

The extension to three-dimensional imaging with CT, as opposed to two-dimensional

mammography, has the advantage that more information may be obtained over a large field of

view and from different imaging planes. Also, imaging is more comfortable for the patient,

since it does not require tissue compression and can be captured with the patient lying down.

However, initial results for CT imaging of lymph nodes in the axilla showed 50% to 82%

sensitivity and 75% to 94% specificity for detection of nodal involvement [53, 54]. This high

variability in accuracy can be attributed to the criterion for differentiation being again only

size [55].

Most recent studies show that PET is more sensitive in imaging the axilla for nodal disease,

with one study reporting a sensitivity of 90% overall, and up to 100% for patients in the T1

(largest) tumour size category [56]. Contrast is based on uptake of radiolabelled glucose

analogues, since there is a higher molecular uptake in fast metabolising cancer cells. However,

the specificity of PET is highly variable (and can be <50%), especially for micro-metastatic

deposits in small lymph nodes, and so the potential for staging the axilla with this modality is

limited [56, 57].

Ultrasound imaging, on the other hand, has demonstrated greater potential for pre-

operative imaging of axillary nodes, especially when combined with FNAC [58, 59] for which

it is becoming an important adjunct clinical imaging modality. In a summary of previous

studies of ultrasound imaging of the axilla, Alvarez et al. found that the sensitivity for

detecting pathological nodes with axillary sonography varied from 44% to 90% depending on

several factors including palpability of the nodes and the number of criterion used. Specificity

varied from 75% to 100% when implemented with FNAC [59]. Standard US assessment is

based on shape and internal morphological appearance, rather than size, to predict lymph

node status. Healthy nodes typically have a thin hypoechogenic cortex and a central echogenic

hilum with a thin border between (Figure 2-5). The length to width ratio (l/w) is an indicator

of node shape and is usually >2 for healthy nodes. A pathological node, on the other hand,

usually exhibits a rounded appearance due to eccentric thickening of the cortex, and l/w is

usually less than 1.5. Absence of the hilum is also highly suggestive of metastasis. Recent

enhancements in US imaging that use colour Doppler flow to image vasculature as an

additional contrast mechanism, have increased the sensitivity and specificity of conventional


Figure 2-5. Ultrasound images of a healthy and abnormal lymph node. (a) Healthy lymph node with a thin cortex (arrowheads) and a large fatty hilum (arrows). Normal nodes are commonly 5 mm or less in length with a length to width ratio of >2. (b) Abnormal lymph node with a uniformly thickened cortex (arrowheads). The thickness of the cortex exceeds that of the fatty hilum so that the length to width ratio is <1.5. The length of malignant nodes is usually greater than 5 mm. Originally published in [48]. © 2007 Radiological Society of North America.

US imaging. A vascular network within the lymph node cortex, as opposed to at the hilum

only, is indicative of cancer growth [50]. Ultrasound is perhaps the most operator-dependent

imaging modality but its value lies in the fact that if suspected metastases are detected, FNAC

can be performed immediately to make a diagnosis and confirm the patient for an axillary

dissection without the need for SLNB.

Dynamic, contrast-enhanced MRI is a more recently investigated modality for detecting

lymph node status and results have shown relatively high sensitivity and specificity.

Assessment with MRI is based on similar criteria to those used in ultrasound, including the

diameter of the short axis of the lymph node and the signal intensity at the hilum and also the

contrast enhancement profile. The earliest study of axillary node assessment with MRI,

reported 90% and 82% sensitivity and specificity, respectively [60]. Recently, advances in MRI

image enhancement, particularly the use of iron oxide nanoparticles and dendrimers to

enhance contrast, have increased the specificity up to 97% [61]. However, the limitations of

MRI, namely, the high cost of installation, the high cost to the patient and artefacts due to

blood flow, have meant this modality has not been widely adopted for pre-operative axillary

imaging and is unlikely to be so in the future.

Overall, these methods have not been sufficiently accurate to be used routinely. One major

reason for this, as already mentioned, is that micro-metastases are difficult to accurately

sample and this leads to low sensitivity. Their utility has relied, rather, on their ability to

identify gross macro-metastatic nodal disease pre-operatively, and identify those patients

suitable for axillary node dissection.

2.2.3 Novel techniques in optical diagnosis

Optical imaging and spectroscopy techniques are fairly new in their development and

application for in situ assessment of soft tissues. Nonetheless, they have great potential for

2.3 Background of optical coherence tomography 19

detecting the signature of cancer due to their high sensitivity, high-resolution, low toxicity and

potential for real-time assessment. Diagnosing disease using optical modalities, termed ‘optical

biopsy’, is based on detecting variations in a tissue’s morphology or biochemical composition

(which is related to tissue function) by measuring the scattering, absorption or emission

processes of light. The measurement is frequently mediated by optical fibres with the optical

probe placed on or near the surface of the tissue. Optical methods have been the subject of

intense research for the detection of cancer and identification of different tissue types during

surgical procedures. The most notable optical techniques include: diffuse optical imaging [7,

62-64]; optical spectroscopies, including fluorescence spectroscopy, elastic-scattering

spectroscopy [65, 66] and Raman spectroscopy [67, 68]; near-infrared fluorescence imaging

[69-71]; fluorescence microscopy, including confocal and multi-photon microscopy [72-76];

and, more recently, optical coherence tomography (OCT) [9, 10, 77, 78], which forms the

focus of the investigations in this thesis and which we discuss further in the following section.

Of those techniques mentioned above, optical spectroscopy has demonstrated the greatest

potential for clinical use. In a review by Bigio and Mourant [79], studies of elastic-scattering

spectroscopy were identified as potential tools for in vivo detection of cancer in the bladder

and gastrointestinal tract, with sensitivities ranging from 60% to 100% and specificities from

80% to 97%. Preliminary research of elastic-scattering spectroscopy for breast cancer detection

and probe guidance during brain surgery were also highlighted. So too were clinical studies of

fluorescence spectroscopy for diagnosis of disease in the oesophagus, cervix and colon, and

preliminary studies of Raman spectroscopy for accurate optical biopsy of cancer in skin, breast

and cervical tissue. The sensitivity and specificity of fluorescence spectroscopy were found

over several studies to range from 83% to 100% and 73% to 98%, respectively.

In summary, optical diagnostic techniques offer the possibilities of higher resolution,

higher sensitivity and faster detection than current clinical imaging modalities. These benefits

could also provide improved disease management by enabling earlier diagnosis and immediate

treatment. The recognised success of these techniques in a range of clinical research areas has

helped drive their progression towards potential clinical applications over the past decade in

particular.

2.3 Background of optical coherence tomography

OCT is a high-resolution, three-dimensional optical imaging modality that has shown great

potential for disease diagnosis in a range of applications. OCT measures elastically


backscattered light to form structural images of tissue, analogous to how ultrasound imaging

measures the reflection of sound.

Since the velocity of light is much faster than that of sound, its reflection cannot be

measured by direct detection based on time-of-flight. Rather, OCT uses interferometric

detection to indirectly measure the location of the reflection by correlating two optical waves,

one reflected from the tissue sample and one from a reference mirror.

The general configuration of an OCT system is that of a Michelson interferometer, which

consists of a broadband light source, a coupler leading to sample and reference arms and a

detector (see Figure 2-9). Depth scans of sample reflectivity are built up either by rapidly

changing the length of the reference path or by performing an inverse Fourier-transform on

the optical spectrum of the interfered waves (these two methods of depth-scanning are

discussed in Section 2.3.1 with further detail given in references [10, 77] and [80]). Cross-

sectional and three-dimensional (3-D) images of tissue are then synthesised from laterally

adjacent depth scans.

The source wavelengths used in OCT are in the near-IR range, in the so-called ‘diagnostic

window’, where absorption due to tissue chromophores is low. In this wavelength range,

scattering dominates over absorption allowing light to penetrate sufficiently to form an image,

usually to a maximum depth of a few millimetres. Figure 2-6 shows some typical components

of tissue that contribute to absorption[81-83]. In the figure, the absorbance of water has been

multiplied by 500 to better visualise it within the plot. Note the high absorbance of water on

the long-wavelength side, and of melanin and deoxyhaemoglobin on the short-wavelength side

of the diagnostic window.

Figure 2-6. Diagram of the optical ‘diagnostic window’ showing the spectral absorption for a range of tissue chromophores. The diagnostic window exists in the near-infrared wavelengths (800–1300 nm), where light attenuation/extinction due to absorption is relatively low. The absorbance of water has been multiplied by 500 to enable its visualisation on the scale of the plot. The arrows indicate typical source wavelengths used for OCT. Composite figure adapted from data in [81] [82] [83].


Light penetration is determined by scattering and absorption. Scattering occurs as a result of

the heterogeneous components that make up tissues including cells, organelles and fibres of

collagen. In general, scattering decreases monotonically with increasing wavelength [84]. The

most common wavelengths for OCT are therefore 1300 nm and 840 nm: although deeper

penetration is achieved at 1300 nm, axial resolution is better at 840 nm (see Equation (2.1)).

The axial resolution of OCT is generally higher than that of standard clinical imaging

systems and approaches that of confocal microscopy as depicted in Figure 2-7. Compared to

conventional microscopy, a major advantage of OCT is its sensitivity, which can reach the

physical limit set by photon shot noise; typically -90 to -110 dB below a perfect reflector. Such

high sensitivity enables penetration up to 2 mm in some tissues, compared to several hundred

microns at best for confocal microscopy. High resolution is also maintained over this depth,

since, unlike confocal microscopy, axial resolution in OCT is decoupled from the numerical

aperture (NA) of the optical system, so that accessing greater imaging depths (through the use

of a low NA objective) does not come at the expense of limited axial resolution [77] (also

illustrated in Figure 2-8). Typical OCT axial resolutions are on the order of 5–10 µm, but

ultrahigh resolutions around 1-2 µm have also been achieved [85]. The low penetration depth

of OCT compared to current medical imaging modalities is a limitation of the technology,

but this may be overcome by the use of compact endoscopic probes and imaging needles that

can be directed to the site of interest. Such probes are now realisable with the development of

new laser and fibre-optic technologies. Advances in detector and acquisition technology are

Figure 2-7. Axial resolution and imaging depth of OCT compared with other imaging modalities. OCT imaging resolution and penetration depth sit at the boundary between potential in situ medical imaging and ex vivo microscopic imaging modalities. The ‘pendulum’ length represents imaging depth, and the ‘sphere’ size represents resolution. From: T. Hillman, “Comparison of OCT with other imaging modalities”, 2009. [Online]. Available: http://obel.ee.uwa.edu.au/research /oct/intro. [Accessed 20 December 2012].


also enabling video-rate, 3-D OCT imaging. Of particular interest for clinical use is functional

OCT imaging, including polarisation-sensitive OCT, Doppler OCT, OCT elastography and

wavelength-dependent OCT. These extensions possess enormous potential for improving

OCT contrast for differentiation of tissue pathology. Overall, OCT is an attractive option for

optical biopsy because it permits real-time, in situ imaging of tissue microstructure with high-

resolution, enabling morphological differences between normal and neoplastic tissues to be

visualised without the need for excision or the addition of exogenous contrast agents. It is also

highly portable, relatively easy to implement and low-cost, making it highly practical for intra-

operative use.

2.3.1 Basic principles of OCT

OCT depth scanning can be achieved using one of two techniques: time-domain low-

coherence interferometry or spectral interferometry. The first implementations of OCT were

based on time-domain OCT (TDOCT) and employed optical low-coherence reflectometry

to generate depth- or A-scans [9, 86]. A standard TDOCT system uses a low temporal

coherence (broadband) light source and a splitter that sends the light into two paths – one

terminating at the sample and the other at a scanning reference mirror as shown in Figure 2-9.

Interference fringes occur only if the (round-trip) optical path lengths of the reference and

sample beams coincide within the coherence length of the source. Therefore, a depth scan is

acquired by scanning the reference mirror and recording the envelope of the generated

interference waveform.

Frequency-domain OCT (FDOCT) is a more recent technique that employs spectral

interferometric techniques to achieve depth sectioning with much faster speeds than

TDOCT [10, 80, 87]. High acquisition speed is possible because the depth scan is generated

by performing an inverse fast Fourier transform (FFT) on the cross spectral density obtained

by interference of the reference and sample beams. Therefore the entire depth scan is obtained

in one computation avoiding the need for the slow scanning reference mirror. The other

major advantage of FDOCT is that it is possible to achieve superior sensitivity with this

technique. This topic has been treated extensively in the literature [80, 88]. Significantly,

unlike TDOCT, the signal-to-noise ratio (SNR) of spectral techniques is independent of the

source optical bandwidth.

There are two distinct methods that have been employed to obtain the spectral signal in

FDOCT. The first uses rapid wavelength tuning of a narrowband source to acquire time-

encoded wavelength information and is called swept-source OCT (SSOCT). The second


method uses a spectrometer to analyse the space-encoded spectral content of the signal and is

called spectral radar or spectral-domain OCT (SDOCT).

For both TDOCT and FDOCT, axial resolution z is given by the full width at half-

maximum (FWHM) of the OCT axial response. Assuming a Gaussian source with a FWHM

bandwidth of , and mean wavelength the axial resolution is [77],

z = 2ln22

,

(2.1)

Transverse resolution is decoupled from axial resolution and is entirely dependent on the

sample arm optics as demonstrated in Figure 2-8. It is given by the FWHM diameter of the

intensity distribution of the sample beam at its waist, with W 0 the 1/e2 diameter of the

sample beam at its waist. For a beam with a 1/e2 waist diameter of W 0 ' incident upon a lens

with focal length f the transverse resolution x, is given by [89],

x = 2 ln2W0 =2 ln2 f

W0 '. (2.2)

The Rayleigh range zR, describes the useful depth-of-field of imaging, where the width of the

beam is within a factor zR times its minimum value. It is given by,

2zR =2 W0

2

. (2.3)

Therefore, a focused beam with a low numerical aperture (NA ≈ / ( W 0) [89]) is usually

desired to achieve reasonable resolution over a large depth-of-field.

Figure 2-8. The OCT sample optics determines transverse resolution and axial depth of field. The coherence gate determines axial resolution. Adapted from: J. G. Fujimoto, “Optical coherence tomography: Introduction”, in Handbook of Optical Coherence Tomography, B. E. Bouma, G. Tearney, eds. New York: Marcel Dekker, Inc., 2001, pp. 1-40.

The final parameter of OCT image quality is sensitivity. In the shot-noise limit, the

sensitivity is proportional to the optical power incident on the sample and inversely

proportional to the A-scan rate and can be calculated by [77],


S =

hc1( )PS PR

PS + PR( )B, (2.4)

where is the quantum efficiency of the detector, h is Plank’s constant, c is the speed of

light, is the intensity splitting ratio of the interferometer splitter (assuming no loss), PS

and PR are the optical powers incident on the sample and reference mirror, respectively and

B is the electronic detection bandwidth, which is a function of the depth scan velocity in

TDOCT, and of the acquisition rate in FDOCT (see references [10] and [90] for more detail

on bandwidth and sensitivity in OCT).

The investigations in this thesis are based on results from three OCT systems. These are:

an in-house TDOCT benchtop system operating at 1310 nm; a fast-scanning commercial

SSOCT system [91] (Thorlabs, USA) operating at 1325 nm; and a second SSOCT system at

1310 nm built during this thesis as a portable system optimised for OCT needle imaging. The

source centre wavelengths for these systems are all around 1300 nm so that they will achieve

high penetration depth when used to image whole lymph nodes. A description of these

systems is given in the following section.

2.3.2 OCT systems used in this thesis

TDOCT

Figure 2-9 shows a schematic of the fibre-based TDOCT benchtop system used for the first

investigation of this thesis (Chapter 3). The splitter/coupler sends polarised light from the

superluminescent diode light source (DenseLight SLED DL-BD9-CS31159A) into the

reference and sample arms. The reference arm is terminated by a frequency-domain optical

delay line (FD-ODL) that rapidly scans the axial path length of the reference arm to achieve

depth sectioning at the sample. It also enables compensation of the dispersion mismatch

between the two arms [92]. Light in the sample arm is directed onto the sample by a weakly

focusing (0.07 NA) triplet lens. The beam is scanned laterally across the lens using a

galvanometer-scanning mirror placed at the back-focus of the lens to accomplish planar

imaging at the sample. Scanning in the third dimension is achieved by translating the sample

on a translation stage.

A circulator at the source arm of the coupler enables balanced detection to reduce

intensity noise. This is achieved by a dual-balanced photodetector (bandwidth = 100 Hz to

3 MHz), which rejects the common-mode DC component of the outputs from the circulator

and the coupler. The dual-balanced photo-detected difference signal is bandpass filtered and


logarithmically demodulated in hardware before analog-to-digital conversion and acquisition

(PCI-6111E, National Instruments, Austin, Texas). Cross-sectional images from this system

were displayed in real-time in Labview using a custom user interface. This system has also been

described previously [93, 94]. A summary of its specifications is given in Table 2-1.

Figure 2-9. Schematic diagram of the fibre-based TDOCT system: SLED, superluminescent light emitting diode; PC, polarisation controller; VOA, variable optical attenuator; FD-ODL, frequency-domain optical delay line; PD, photodetector.

Table 2-1. Theoretical and measured performance parameters for the 1310 nm TDOCT system

Parameter Theoretical Measured

Centre wavelength ( 0 ) 1310 nm 1320 nm

Source bandwidth ( B3dB ) 130 nm (100 nm before upgrade)

154 nm (100 nm before upgrade)

Average source power ( P0 ) 15 mW 21 mW

Axial resolution in air ( zFWHM, intensity )

6.3 µm (8.9 µm before upgrade)

8.7 µm (11.2 µm before upgrade)

Transverse resolution ( xFWHM ) 6.6 µm 6.4 µm

Depth of field ( 2zR ) 127 µm 260 µm

Sensitivity ( S ) -112 dB (shot noise limited) -108 dB

A-scan rate 500 Hz 500 Hz

Depth scan range in air ( Z ) 2.26 mm 2.3 mm

Transverse scan range* ( X , Y )

— 3.7 mm × 6.0 mm

Typical sampling density (pixel size) ( Z , X ,Y ) — 2.1 µm × 3.4 µm × 5.9 µm

Power at sample ( PS ) — 2.9 mW *Determined directly by the settings of the galvanometer scanner and translation stage at the sample


The advantages of TDOCT are its relative low complexity and ease of implementation,

with components for this system being readily attainable (the broadband source and

associated passive components of the interferometer were upgraded during this thesis). Its

main disadvantage is its very slow imaging rate of 2 seconds per B-scan with 3-D scans taking

up to 35 minutes, which is not feasible for clinical studies.

SSOCT

Schematics of the Thorlabs 1325 nm SSOCT system (OCS1300SS) and the custom portable

1310 nm SSOCT system used in this thesis are shown in Figure 2-10 and Figure 2-11,

respectively. Both systems use a frequency clock (Mach-Zehnder interferometer (MZI) clock)

so that sampling at the interferometer output is linear in k-space (wavenumber), since the

Fourier transform pair is wavenumber and group delay. Both systems employ balanced

detection, although this is slightly different in the custom-built system, which was made for

maximum flexibility for using OCT needle probes in both dual-arm Michelson configuration

and common path Fizeau configuration. Thus, the balancing arm in this system is provided by

the output of a second splitter at the source, which takes power from the source to balance its

intensity noise as shown in Figure 2-11.

Scanning at the sample arm of the Thorlabs system was performed by an XY galvo-

canning mirror pair enclosed in a detachable handheld scan head, which included a visible

CCD camera to view the scan area. The mirrors translate the beam over a maximum area of

10 × 10 mm at the sample. All data acquisition and processing was performed via the

integrated software package and PC provided by Thorlabs. The system achieved high imaging

rates; however, the axial and lateral resolutions were lower than the TDOCT system. Also it

Figure 2-10. Schematic diagram of the Thorlabs SSOCT system: SS, swept laser source; FC, fibre coupler; PC, polarisation controller; CIR, circulator; C, collimator; AP, adjustable pinhole variable attenuator; M, mirror; BD, balanced detector; DAQ, data acquisition card; SD, XY scanners driver; CCD, camera; OBJ, objective. From Thorlabs OCS1300SS Swept Source OCT System User Guide.


Table 2-2. Theoretical and measured performance parameters for the Thorlabs SSOCT system.

Parameter Provided by supplier Measured


Source bandwidth ( B3dB ) 100 nm 100 nm

Average source power ( P0 ) 10 mW 13.6 mW

Axial resolution in air ( zFWHM, intensity )

12 μm 17 µm

Transverse resolution ( xFWHM ) 15 μm 16 µm

Rayleigh range ( zR ) 430 μm 416 µm


A-scan rate (sweep rate) 16 kHz 16 kHz

Depth scan range in air ( Z ) 3.0 mm 2.8 mm

Transverse scan range ( X , Y ) 10 mm × 10 mm (max) 4 mm × 4 mm (Typical)

Sampling density (pixel size) ( Z , X ,Y )

— 5.3μm × 5μm × 5μm (Typical)

Depth scan range in air ( Z ) 5.0 mW 4.7 mW

was difficult to adjust and optimise the processing and image acquisition parameters for our

application. This also meant that replacing the sample arm with a needle sample arm would

not be possible: for example, the scanning control was integrated into the manufacturers

software assuming XY galvo scanning, but this could not be altered to allow needle scanning

control. For these reasons we built the second SSOCT scanner shown in Figure 2-11, which

was also more compact and portable to be used for OCT needle imaging at the hospital.

Further details of this system along with the needle sample arm scanning are given in Chapter

5. A summary of the specifications for each system is given in Table 2-2 and Table 2-3.

The major advantage of FDOCT over TDOCT is that imaging is much faster (Thorlabs

system scan rate = 25 Hz B-scan rate). Also, FDOCT is known to have a sensitivity advantage

of up to 20 dB over TDOCT [80]. The reference arm is also stationary and more compact

than TDOCT making it easy to align and portable. In the case of SSOCT, both reference arm

alignment and signal detection is simple and easily made portable. One disadvantage of

SSOCT is that, due to limitations in the bandwidth of currently available swept sources, the

axial resolution is limited; although advances in source technology, such as Fourier domain

mode-locked wavelength swept lasers and polygon scanners are leading to improvements in

source bandwidth and speed [95]. Also dispersion compensation is slightly more difficult and


Figure 2-11. Schematic diagram of the custom-built portable SSOCT system: SS, swept source; VOA, variable optical attenuator; DAQ, data acquisition card; PD, photodetector; tr, sweep trigger; clk, k-linear source clock.

Table 2-3. Theoretical and measured performance parameters for the 1310 nm custom-built needle SSOCT system.

Parameter Theoretical Measured


Source bandwidth ( B3dB ) 100 nm 100 nm

Average source power ( P0 ) 18 mW (min) 41.5 mW

Axial resolution in air ( zFWHM, intensity ) 13 μm 13 μm

Transverse resolution ( xFWHM )

— *

Rayleigh range ( zR ) — *


A-scan rate (sweep rate) 50 kHz 50 kHz

Depth scan range in air ( Z ) 5 mm 4.7 mm

Transverse scan range ( X , Y ) — 360° × 6 mm (Typical)

Sampling density (pixel size) ( Z , X ,Y ) — 4.9μm × 0.5° × 5μm (Typical)

Depth scan range in air ( Z ) — 18 mW *The transverse resolution and Rayleigh range are determined by the optics of the needle probe used in each experiment (see Chapter 5 for resolutions of the needle probes used in this thesis )

generally needs to be implemented numerically in software after acquisition. Possible

improvements to the OCT technology, as well as the use of functional and multimodal

extension to OCT imaging, are briefly discussed in Chapter 6. The current work focuses on

the use of standard OCT and in the next chapter we present the application of TDOCT for

imaging of excised healthy and malignant human axillary lymph nodes.

optical coherence tomography for optical biopsy of

Documents