1

MEMS BASED OPTICAL COHERENCE TOMOGRAPHY IMAGING

By

JINGJING SUN

A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT

OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY

UNIVERSITY OF FLORIDA

2012

2

© 2012 Jingjing Sun

3

To my parents, Xueyi Sun and Xianghui Wang

4

ACKNOWLEDGMENTS

I would like to express my deepest gratitude to my advisor, Dr. Huikai Xie, for his

guidance and support through the course of my study and research. His profound

knowledge and extensive experience in the field of optical MEMS helped me overcome

many obstacles. My research would not be successful without his encouragement and

patience. I would also like to thank my committee members, Dr. Henry Zmuda, Dr. Peter

Zory, and Dr. Malisa Sarntinoranont for their valuable suggestions for my work. I

benefited a lot from Dr. Zmuda’s optics class -- his well structured lectures provided me

a better understanding of the field of optics. I would like to thank Dr. Zory for generously

providing his equipments to our lab, a lot of my experiments were conducted with Dr.

Zory’s equipments. And I would like to thank Dr. Sarntinoranont for her help in the

refractive index measurement project; her useful suggestions really helped improved

the quality of my research. In addition, I would like to thank Dr. Brian Sorg for providing

resources for the animal experiments, Dr. Huabei Jiang for collaborating with our lab on

the photoacoustic research, and Dr. Antonio Pozzi for his collaboration on the PS OCT

study.

I have received a lot of help along the way. I would like to thank Dr. Shuguang

Guo for introducing me to the world of OCT, when I first started working on this project.

He really helped laying a solid foundation for my future research. I would also like to

thank Dr. Lei Wu for providing me with MEMS mirrors to work with, and passing me a lot

of knowledge on MEMS. I thank Dr. Sung Jin Lee for his help in the refractive index

measurement project; Dr. Lee has given me a lot of help and insight on the experiment

and paper writing process. I would also like to thank Lei Xi and Dr. Yiping Zhu, with

whom I worked on the OCT and photoacoustic project. I thank Dr. Se-Woon Choe for

5

helping me with the animal experiments. I would also like to thank Can Duan, who has

helped me a lot with experiments since she joined our group. My thanks also go out to

everyone in BML and IMG for their useful discussions and friendships.

Last but not least, I would like to thank my loving parents and my boyfriend for

their love and support.

6

TABLE OF CONTENTS page

ACKNOWLEDGMENTS .................................................................................................. 4

LIST OF TABLES ............................................................................................................ 9

LIST OF FIGURES ........................................................................................................ 10

ABSTRACT ................................................................................................................... 14

CHAPTER

1 INTRODUCTION .................................................................................................... 16

Traditional Cancer Detection Methods and Their Limitations .................................. 17 Biomedical Optical Imaging .................................................................................... 17

Confocal Laser Scanning Microscopy .............................................................. 18 Nonlinear Optical Microscopy ........................................................................... 19 Optical Coherence Tomography ....................................................................... 20

Limitation in OCT Imaging for Endoscopic Application ........................................... 21 Introduction to MEMS Technology .......................................................................... 21

MEMS Based Endoscopic OCT Imaging ................................................................ 23 Endoscopic Probes Overview ........................................................................... 23 MEMS OCT Endoscopic Probes Overview ...................................................... 24

Electrostatic MEMS Endoscopes ............................................................... 25

Electromagnetic MEMS Endoscopes ......................................................... 28

Piezoelectric MEMS Endoscopes .............................................................. 31 Electrothermal MEMS Endoscopes ........................................................... 32

Research Goals ...................................................................................................... 35 Thesis Outline ......................................................................................................... 36

2 OPTICAL COHERENCE TOMOGRAPHY .............................................................. 37

Low Coherence Interferometry ............................................................................... 37 Time Domain OCT System ..................................................................................... 40 Frequency Domain OCT ......................................................................................... 43

Principle of Frequency Domain OCT ................................................................ 44 Configurations of FD-OCT ................................................................................ 46

SNR Advantage of FD-OCT ............................................................................. 48 Functional OCT ....................................................................................................... 49

Doppler OCT .................................................................................................... 50 Polarization Sensitive OCT ............................................................................... 52

Time Domain OCT System in BML ......................................................................... 57 Summary ................................................................................................................ 61

7

3 ELECTROTHERMALLY ACTUATED MEMS MIRRORS ........................................ 62

Electrothermal Actuation ......................................................................................... 62 1st and 2nd Generations Electrothermal MEMS Mirrors ........................................... 63

LSF-LVD (3rd) and FDSB (4th) Electrothermal MEMS Mirrors ................................. 65 LSF-LVD MEMS Mirror..................................................................................... 65 Light Weight Large Aperture LSF-LVD MEMS Mirror ....................................... 69 FDSB MEMS Mirror .......................................................................................... 70

Summary ................................................................................................................ 73

4 OCT ENDOSCOPIC IMAGING BASED ON MEMS ............................................... 75

MEMS Endoscopic Probe Design ........................................................................... 75 Probe Optical Design ....................................................................................... 76

Probe Mechanical Design ................................................................................. 77 In Vivo Experiments ................................................................................................ 83

Issues with Probe Imaging ...................................................................................... 86

Effect of the Distance between the Fiber and the GRIN Lens .......................... 89 Effect of Mirror Curvature ................................................................................. 91

Effect of FEP Tubing ........................................................................................ 93 Effect of Non-Telecentric Scan ......................................................................... 96

Summary ................................................................................................................ 99

5 MEMS BASED FREE-SPACE TISSUE IMAGING — REFRACTIVE INDEX MEASUREMENT .................................................................................................. 100

RI Measurement Background and Motivation ....................................................... 100 Free Space OCT System ...................................................................................... 102

Sample Preparation .............................................................................................. 103 Method and Process ............................................................................................. 104 Results .................................................................................................................. 108

Discussion and Conclusion ................................................................................... 110

6 MEMS FUNCTIONAL OCT IMAGING (PS, DOPPLER) ....................................... 114

PS OCT System ................................................................................................... 114 PS OCT System Configuration ....................................................................... 114 Ex Vivo PS OCT Imaging Experiment ............................................................ 118

Doppler OCT System ............................................................................................ 120

Doppler OCT Principle.................................................................................... 120 In Vivo Doppler OCT Imaging Experiment...................................................... 121

Summary .............................................................................................................. 122

7 MEMS OCT/PHOTOACOUSTIC IMAGING .......................................................... 123

Photoacoustic Imaging System ............................................................................. 123 Principle and Background............................................................................... 123 MEMS Based Photoacoustic Probe ............................................................... 124

8

Photoacoustic System .................................................................................... 125

Photoacoustic Experiments ............................................................................ 126 MEMS Probe for Both OCT and Photoacoustic Imaging ...................................... 127

MEMS Probe Design ...................................................................................... 127 Ex Vivo Experiment ........................................................................................ 128

Summary .............................................................................................................. 129

8 CONCLUSION ...................................................................................................... 130

Summary of Work ................................................................................................. 130

Future Work .......................................................................................................... 131

APPENDIX: PUBLICATIONS ...................................................................................... 133

LIST OF REFERENCES ............................................................................................. 135

BIOGRAPHICAL SKETCH .......................................................................................... 150

9

LIST OF TABLES

Table page 3-1 Optimized structure parameters of FDSB actuator ............................................. 72

10

LIST OF FIGURES

Figure page 1-1 Confocal imaging system schmatic .................................................................... 18

1-2 A simplified schematic for TPEF and SHG imaging set-up ................................. 19

1-3 OCT system schematic ...................................................................................... 20

1-4 Existing MEMS devices. ..................................................................................... 22

1-5 Commercially available endoscopes [42][43]. .................................................... 24

1-6 MEMS probe schematic.. ................................................................................... 25

1-7 Electrostatic MEMS OCT from UC-Irvine ........................................................... 27

1-8 Electrostatic MEMS OCT from MIT.. .................................................................. 28

1-9 Electromagnetic MEMS OCT reported by Kim et al. [71]. ................................... 30

1-10 Electromagnetic MEMS OCT from Yamagata Research Institute of Technology, Japan. [72]. .................................................................................... 30

1-11 Piezoelectric MEMS OCT reported by Gilchrist et al. [79]. ................................. 32

1-12 Electrothermal MEMS OCT reported by Pan et al. ............................................. 33

1-13 Electrothermal MEMS OCT reported by Xu et al. ............................................... 34

2-1 Michelson interferometry.. .................................................................................. 39

2-2 Gaussian distribution broadband light source interference pattern. .................... 39

2-3 OCT sample signal. ............................................................................................ 40

2-4 RSOD configurations.. ........................................................................................ 42

2-5 Examples of applications of time domain OCT system. ...................................... 43

2-6 Example application of FD-OCT system ............................................................. 44

2-7 FD-OCT signal detection schematic. .................................................................. 46

2-8 FD-OCT configurations.. ..................................................................................... 47

2-9 Doppler OCT sample beam ................................................................................ 50

11

2-10 Consecutive A-Line scans in Doppler OCT ........................................................ 52

2-11 Free space PSOCT system setup. ..................................................................... 54

2-12 OCT system schematic. ..................................................................................... 57

2-13 RSOD configuration ............................................................................................ 58

2-14 Ray trace of RSOD with no center misalignment. ............................................... 58

3-1 Eletrothermal bimorph structure. ........................................................................ 63

3-2 1st generation MEMS mirrors.. ............................................................................ 64

3-3 2nd generation: lateral shift free (LSF) MEMS mirror design. .............................. 65

3-4 LSF-LVD eletrothermal actuator and mirror design ............................................ 66

3-5 Electron micrographs of the 2D MEMS mirror.. .................................................. 67

3-6 MEMS mirror experimental results. .................................................................... 68

3-7 MEMS mirror used for lateral scan.. ................................................................... 69

3-8 FDSB actuator design. ....................................................................................... 71

3-9 FDSB MEMS mirror.. .......................................................................................... 71

3-10 Characterization of the FDSB mirror.. ................................................................. 72

4-1 Typical MEMS probe design. .............................................................................. 76

4-2 Optical components of the probe. ....................................................................... 76

4-3 Mechanical dimensions of the probe .................................................................. 78

4-4 1st generation probe design. ............................................................................... 79

4-5 2nd generation probe design.. ............................................................................. 80

4-6 3rd generation probe design.. .............................................................................. 81

4-7 DFSB probe design.. .......................................................................................... 82

4-8 DFSB probe design.. .......................................................................................... 83

4-9 In vivo OCT images of mouse tongue. ............................................................... 84

4-10 In vivo images of mouse ear.. ............................................................................. 85

12

4-11 In vivo imaging results.. ...................................................................................... 86

4-12 Demonstration of issues focused on in this study. .............................................. 87

4-13 Code V simulation.. ............................................................................................ 88

4-14 Optical quality testing experiments. .................................................................... 89

4-15 Effects of the distance between the fiber and the GRIN lens ............................. 90

4-16 Code V simulation for effect of mirror curvature. ................................................ 91

4-17 Effects of the MEMS mirror radius of curvature based on simulation. ................ 93

4-18 Comparison of simulation and experimental results. .......................................... 93

4-19 Effects of FEP tubing on optical beam from the GRIN lens. ............................... 95

4-20 Effects of FEP tubing on the system. .................................................................. 96

4-21 Demonstration of the non-telecentric scan effect. ............................................... 98

4-22 Correction for curved image display.. ................................................................. 98

4-23 Non-telecentric scan effect in radial direction. .................................................... 99

5-1 OCT setup for refractive index measurement. .................................................. 103

5-2 Anatomical regions of rat brain tissue slices tested. ......................................... 105

5-3 Measurement process.. .................................................................................... 106

5-4 Measurement of optical and physical thickness of various anatomical regions in rat brain tissue slices.. .................................................................................. 107

5-5 OCT images of brain tissue slices under compression. .................................... 108

5-6 Measured RI in various anatomical regions in brain and aCSF. ....................... 109

5-7 RI in rat brain tissue slices under uniform compression. .................................. 110

6-1 Schematic of PS OCT system. ......................................................................... 115

6-2 PS OCT imaging of canine meniscus. .............................................................. 119

6-3 PS OCT imaging results.. ................................................................................. 119

6-4 In vivo Doppler imaging. ................................................................................... 121

7-1 Endoscopic photoacoustic probe. ..................................................................... 125

13

7-2 Photoacoustic system. ...................................................................................... 126

7-3 Ex vivo photoacoustic experiment with pencil lead.. ......................................... 127

7-4 In vivo photoacoustic experiment. .................................................................... 127

7-5 Probe for OCT and PA imaging. ....................................................................... 128

7-6 OCT images of mouse ear................................................................................ 129

14

Abstract of Dissertation Presented to the Graduate School of the University of Florida in Partial Fulfillment of the Requirements for the Degree of Doctor of Philosophy

MEMS BASED OPTICAL COHERENCE TOMOGRAPHY IMAGING

By

Jingjing Sun

May 2012

Chair: Huikai Xie Major: Electrical and Computer Engineering

Optical coherence tomography (OCT) can provide subsurface cross-sectional

information of tissue samples with high resolutions of 1-15 µm. Microelectromechanical

system (MEMS) mirrors are small and can perform fast optical scans; they are suitable

for realizing lateral optical scans in OCT systems. In this thesis, MEMS technology is

applied to enable endoscopic OCT imaging. Two generations of MEMS mirrors have

been used. The MEMS mirrors are 2 mm x 2 mm and 1.7 mm x 1.55 mm respectively in

size. Four generations of MEMS-OCT endoscopic probes have been developed; the

diameters of the probes are reduced from 5 mm in the first generation down to 2.5 mm

in the fourth generation. The effects of the positions of the optical components and the

protective tubing on OCT imaging performance have been studied through simulation

and verified experimentally. In vivo imaging experiments were done on mouse tongue,

mouse ear and injected tumor; layered structures of mouse tongue and ear, and

accurate shape of tumor regions were clearly detected.

The same MEMS-OCT system has also been used to measure localized refractive

index of acute rat brain tissue slices; regions including cerebral cortex, thalamus, corpus

callosum, hippocampus and putamen were measured; refractive index in the corpus

15

callosum was found to be ~ 4% higher than the RIs in other regions. Changes in

refractive index with tissue deformation were also measured in the cerebral cortex and

corpus callosum under uniform compression (20-80% strain). For 80% strain, measured

RIs increased nonlinearly by up to 70% and 90% in the cerebral cortex and corpus

callosum respectively. Further, two types of functional OCT have been realized:

polarization sensitive (PS) OCT and Doppler OCT. PS OCT has been successfully

applied on canine meniscus, and tissue birefringence were detected and represented by

their Stokes parameters shown by 2D and 3D images. Rat blood vessels were identified

by Doppler OCT. In addition, MEMS based OCT combined with photoacoustic imaging

was demonstrated.

This is the first detailed study on applying MEMS technology to realize endoscopic

OCT imaging. This work opens up many promising opportunities in early cancer

detection and real-time imaging guided surgery.

16

CHAPTER 1 INTRODUCTION

Cancer has been one of the fatal diseases in the United States. Nearly one in four

deaths are caused by cancer in the U.S. American Cancer Society (ACS) predicts that

there will be 1,638,910 cancer cases in the year 2012, and among them, 577,190

cancer patients are not expected to live [1]. Reducing cancer death rate has become a

global research focus. Early detection and improved treatment have proven to be

effective, as cancer survival rate increased from 49% to 67% over the past 30 years [1].

The development of biomedical optical imaging modalities, especially optical

coherence tomography (OCT) has been a major contribution to early cancer detection;

as it can provide in vivo non-invasive 3D ultrahigh resolution images that allows us to

see cancer or precancerous lesions that were not visible to traditional imaging

techniques, including X-ray, computer tomography (CT) scan, ultrasound and magnetic

resonance imaging (MRI) [2].

Endoscopic imaging is very important, because most cancers occur in internal

organs; yet it is difficult, because fast 2D lateral scan needs to be realized in small

endoscopes that must be able to fit into narrow lumens whose diameters typically are

only a few millimeters. Micro-electro-mechanical system (MEMS) technology is an

enabling technology that makes devices and systems on the scale of micrometers [3]-

[5].Development of MEMS mirrors has shown great potential for their application as

endoscopic optical scanners. This chapter discusses the research motivation, and

remaining challenges to the existing solutions, then our solution will be proposed,

followed by an outline of this thesis.

17

Traditional Cancer Detection Methods and Their Limitations

Surgical biopsy has been the conventional cancer detection method, which suffers

from several drawbacks: first, biopsy is an invasive procedure; tissue samples need to

be cut off from questionable areas of diseased organs from human body for pathological

examinations; second, biopsy results may not be reliable, due to the fact that the testing

samples are randomly selected from the target region; third, the diagnostic cycle is long,

which is caused by time-consuming ex vivo experiments to be conducted. Biomedical

imaging techniques are promising alternatives to biopsy; the imaging procedures are

non-invasive and the diagnosis cycles are shorter. Two key factors for early cancer

detections are: first, the imaging modalities should be able to detect precancerous

lesions, whose sizes are normally around 5 to 10 µm; and second, the imaging depth

should be able to cover the epithelial layer, where most early cancerous morphological

changes occur [6], [7]. Conventional bio-imaging techniques that are already widely

used in clinical diagnoses include X-ray, computer tomography (CT) scan, ultrasound

and magnetic resonance imaging (MRI). Among these methods, CT scan and MRI are

the most popular modalities for tumor imaging. However, they are not suitable for early

cancer detection due to their low resolution, which is on the order of 100 µm. Biomedical

optical imaging modalities, which can realize ultrahigh resolution imaging (around 10

µm) are therefore desirable for this task.

Biomedical Optical Imaging

Confocal laser scanning microscopy (CLSM), nonlinear optical microscopy

(NLOM) and optical coherence tomography (OCT) are among the most popular

biomedical imaging techniques; all three of them are capable of providing ultrahigh

18

resolution 2D and 3D cross-sectional tissue images. This section discusses the

principles of these imaging techniques.

Confocal Laser Scanning Microscopy

CLSM has been widely used for tissue imaging and neuron imaging due to its high

resolution compared to traditional microscopy [8]. Clinical applications in evaluating eye

and bone diseases have also been reported [9], [10]. Ultra high depth resolution of less

than 1 µm in CLSM can be achieved by employing pinholes to reject out-of-focus light,

as shown in Figure 1-1. Only optical signal from the focal plane can reach the detector;

therefore, optical sectioning can be achieved. Imaging depth of CLSM can reach up to

100 µm, which is limited by the scattering effect [8]. 3D image can be reconstructed by

combining depth scan and 2D lateral scan [11]. Depth scan is generated by displacing

the pinhole through various axial positions to change the focal planes; while lateral scan

moves the pinhole in lateral directions to cover desired scanning area. 3D images can

be constructed using pixel–by –pixel signal collection.

Figure 1-1. Confocal imaging system schmatic

Photodetector

Pinhole

Aperture Beam Splitter

Pinhole

Aperture

Focal Plane

Sample Light

Source

19

Nonlinear Optical Microscopy

NLOM uses higher order light-matter interaction to generate optical signals. Based

on nonlinear processes that can occur during light-tissue interaction, there are several

types of NLOM. The most researched ones include two photon–excited fluorescence

laser-scanning microscopy (2PLSM) [12], [13], second harmonic generation (SHG) [14]-

[16] and coherent anti-Stokes Raman scattering (CARS) [17], [18]. NLOM is less

sensitive to scattering, so it can be used to generate larger imaging depth up to 1mm.

Compared to linear light-tissue interaction, the chances of nonlinear interaction is

very low, and only when the incident light is focused in both time domain and spatial

domain, can enough fluorescence optical signal be collected for image reconstruction.

High spatial resolution can therefore be achieved without the usage of a pinhole [13].

Objective lens can be used to focus light spatially, while femtosecond pulsed lasers are

required to focus the optical energy in time domain. 3D capability has also been

demonstrated for NLOM.

Figure 1-2. A simplified schematic for TPEF and SHG imaging set-up

Ti: Sapphire

Laser Prechirp Unit

Iris

filter

PMT

Bandpass

Filter

Dichroic Mirror

Objective Lens

Double-clad Photonic

Crystal Fiber

Scan

Mirror

Sample

20

Optical Coherence Tomography

OCT employs an interferometer structure, and relies on coherence gating to reject

signal back scattered from sample outside the coherence length [2]. OCT system uses

a broadband light source at the near infrared range. This wavelength has relatively large

penetration depth for tissues; depending on the tissue type, about 1 to 3 mm imaging

depth could be achieved. Broadband light source can also effectively reduce the

coherence length; therefore, axial resolution can be as high as 1 to 15 µm. Depth scan

of OCT is realized by moving the reference arm back and forth, when combed with 2D

lateral scan on the sample arm, 3D imaging can be generated.

Figure 1-3. OCT system schematic

As discussed above, all three biomedical optical imaging modalities have the

ability to distinguish precancerous lesions from healthy lesions, and to obtain 3D

images. Yet, only OCT has imaging depth large enough to cover the epithelial layer,

therefore, it is most suitable for noninvasive early cancer detection. In addition, OCT

uses relatively lower incident optical power compared to CLSM and nonlinear imaging

systems, which makes it safer for human tissues imaging.

Light source 2×2 Beam

splitter

Reference mirror

Photo detector

Lateral scanning DAQ PC

21

Limitation in OCT Imaging for Endoscopic Application

OCT, since its introduction in the early 1990s, has been employed extensively in

dermatology and ophthalmology [2], [19]-[25]. It has also been used for imaging internal

organs such as the GI tracks, bladders, and esophagus [26]-[30], [33]-[36]. As

mentioned before, endoscopic tissue imaging is very important, yet it is one of the most

challenging part of OCT imaging. Numerous attempts have been made to tackle this

problem. Rotating a fiber micro-prism module at the proximal end [30], and swinging a

distal fiber tip by a galvanometric plate [33] are two earlier ones. These methods are

simple, but relatively slow. In order to increase lateral scanning speed, one method is to

excite a cantilever fiber to its resonance by a piezoelectric tube [35]. The drawback of

this method is the loss of coupling efficiency due to the fiber tip vibration. To solve this

problem, one way proposed is to use a pair of GRIN lens with proper cut angle at the

ends, transverse scanning is achieved by rotating the two lenses at matching speeds

[36]. Yet, further miniaturization is difficult for this design, because two motors are

needed for rotation. In this work, we propose a solution, which integrates MEMS

scanners with OCT to achieve fast lateral scans in small endoscopic probes.

Introduction to MEMS Technology

Micro-electro-mechanical systems (MEMS) technology makes mechanical or

electro-mechanical devices, structures or systems using techniques of microfabrication.

MEMS devices are small and their sizes range from tens of microns to a few

millimeters. MEMS sensors and actuators have been widely used in many products. For

example, MEMS accelerometers can sense changes in velocity, and have been used in

automotives for airbag deployment to improve drivers’ safety. The same accelerometers

can be adapted to sense the tilting of cell phones, tablets, etc. in order to correct the

22

screen display directions. MEMS gyroscopes can sense the change in tilting angles;

many GPS incorporate MEMS gyroscopes and sometimes accelerometers to improve

their performances. One of the most popular applications of MEMS actuators are optical

scanners. There are two types of optical scanners: digital and analog. Digital

micromirrors can be set to two states: on and off. Texas Instrument’s Digital Micromirror

Devices (DMDs) are digital micromirrors, and they are the core of the Digital Light

Processing (DLP) technology [37]. Now about half of the world’s projectors are using

the DLP chips. Analog micromirrors are more flexible, they can steer light continuously

over a certain range. Applications including barcode scanning, virtual retinal display and

projection display have been reported [38]-[40].

Figure 1-4. Existing MEMS devices. A) Draper tuning fork gyroscope [41], B) TI DMD mirrors[37].

The successful application of MEMS technology in other fields has demonstrated

some great features of MEMS devices, which are desirable for biomedical applications.

First, MEMS is small. MEMS mirrors can be only a few millimeters in diameter, which

makes it ideal for working as a scanning engine in endoscopic biomedical imaging

probes. Second, MEMS is fast. MEMS scan engines can work at hundreds of Hertz or

A B

23

even up to thousands of Hertz during scanning, therefore, real time imaging can be

possible. Third, the cost of MEMS is low. MEMS devices can be mass produced, since

it borrowed fabrication process from the mature semiconductor industry; consequently,

the cost for each device is low. In addition, MEMS has the advantages of low power

consumption and ease of integration. As a result, MEMS mirrors would be the favorable

choice for OCT scan engines.

MEMS Based Endoscopic OCT Imaging

Endoscopic Probes Overview

Endoscopes are tube-like biomedical instruments that are used to look inside

human body or to perform certain surgical procedures. Depending on the body parts

being examined, the types and sizes of endoscopes have different requirements. To

look at body parts that are more easily accessible, for example, one’s ear, nose or

throat, a rigid probe would be more suitable; while to look at organs deeper inside the

body, such as the GI tracks, a flexible probe may be required.

A flexible endoscope normally has three sections: the control section, the

connector section and the insertion tube [43], [44]. The control section is the part held

by the doctor, it normally has angulation controls knobs, valves for suction or air/water,

and the instrument port. Electrical connection of the probe is located at the connection

section; it also has a light guide, an air pipe, a suction port and a safety connecter

mount. The insertion tube is the part that actually goes inside the human body. It

normally has illumination lens, an air/water vent, an instrument channel for suction or

biopsy, and an objective lens for imaging. The diameter of the insertion tube normally

ranges from 2.8 mm to 13.7 mm. OCT endoscopic probes could be designed to be fitted

24

in an insertion tube of the commercially available endoscopes as described, or it could

be inserted into human biopsy channel directly for OCT imaging.

Figure 1-5. Commercially available endoscopes [42][43]. A) rigid probe, B) flexible probe, C) control section of flexible probe, D) connection section of flexible probe, and E) distal end of the insertion tube.

MEMS OCT Endoscopic Probes Overview

The first MEMS based OCT endoscope was introduced by Y. Pan et al. employing

a one-dimensional (1D) electrothermally-actuated MEMS mirror [45]. Two-dimensional

(2D) porcine bladder cross-sectional images were demonstrated. After that, various

forms of MEMS mirrors have been developed as the scanning engine in endoscopic

probes for OCT systems.

Figure 1-6 illustrates the concept of MEMS-based front-view and side-view OCT

probes, in which MEMS mirrors are placed at the distal ends of the probes, and their

angular rotation directs the light and generates lateral scans on the sample. Several

features are highly desirable for this application. Firstly, the footprint of the MEMS

B

C

A

Instrument channel

Objective lens Air/water vent

Light

D E Electrical connection

Light guide

Air pipe Suction port

S-cord connection mount

25

device must be small to fit into small endoscopes; secondly, the mirror aperture must be

large and flat for easy optical alignment and high optical resolution; thirdly, the mirror

must be able to scan large angles to realize large imaging area; fourthly, the driving

voltage must be low to ensure safe use inside human body; and finally, linear control of

the scan should be easily implemented to simplify signal processing and image

interpretation. Single-crystal-silicon (SCS) or silicon-on-insulator (SOI) substrate can

provide large, flat and robust device microstructures; thus they are predominantly used

for making MEMS mirrors for endoscopic applications. There are four actuation

mechanisms employed for generating the scanning motion for MEMS mirrors:

electrostatic, electromagnetic, piezoelectric and eletrothermal actuations. All four

mechanisms have been successfully used for OCT endoscopes, and will be reviewed in

the following sections.

Figure 1-6. MEMS probe schematic. A) front view probe, B) side view probe.

Electrostatic MEMS Endoscopes

Electrostatic actuation has been one of the most popular choices for MEMS

mirrors. Electrostatic actuation is based on electrostatic force which exists between

electrically charged particles. The first MEMS mirror based on this principle was

demonstrated by Petersen in 1980 [46]. By employing a parallel-plate structure, the

maximum rational angle of the micromirror reached ±2º at resonance and 300V driving

A B

26

voltage. This parallel-plate structure has been adopted by many other researchers and

rotation angles as large as ±8º have been achieved with lower driving voltages in the

range of 40V to 200V, but the reported mirror aperture diameters are normally smaller

than 500µm [47]-[54]. What limits this structure from realizing larger scan angle is the

pull-in effect. To overcome this problem, vertical comb drive actuation structures were

employed [55]-[60]. The vertical or angular displacement of the comb drives is

converted into mirror tilting through a torsional beam which supports the mirror plate.

Mirror aperture size as large as 1.5 mm × 1.5 mm has been reported based on this

structure [56]. Both 1D and 2D micromirrors have been realized using this structure, and

the mechanical deflection angle has reached ±5.5º at only 16 V resonance [56] and

±6.2º at 55 V resonance [57]. For 2D mirrors, two gimbals are used to support the mirror

plate in orthogonal directions. To further increase scan angles, Milanovic et al. proposed

a new gimbal-less SOI-based micromirror [58], the static optical deflection for both axes

have increased to ~ ±10° with ~150 V driving voltage.

A series of endoscopic OCT probes employing electrostatic micromirrors have

been reported by Jung and McCormick [61], [62]. Figure 1-7 A shows the SEMs of two

of the MEMS mirrors they reported. The one on the left has a 1 mm × 1 mm mirror

aperture size on a 2.8 mm × 3.3 mm footprint, while the one on the right has a mirror

aperture diameter of 800 µm. The devices employed 2D gimbal-less vertical comb

structures. The mirror plate and actuators were fabricated separately and then bonded

together. Optical scanning angles of the mirrors can reach 20° at resonance when a 100

V driving voltage is applied. Resonant frequencies as high as 2.4 kHz and 1.9 kHz on

two axes have been reported. Electrical connection was done through wire bonding,

27

and injection mold was used for probe body. The diameters of the endoscopic probes

range from 3.9 mm to 5.0 mm. 3D in vivo OCT images of healthy rabbit rectal tissues

have been obtained at 8 frames/sec. The image volume size is 1 × 1 ×1.4 mm3, and the

resolution of the image is 20 × 20 × 10 µm3. Important tissue structures can be clearly

seen here.

Figure 1-7. Electrostatic MEMS OCT from UC-Irvine. A) SEMs of two MEMS mirrors, B) Electrical connection through wire-bonding, C-E) packaged MEMS probe, F) 3D image of rabbit rectal tissue.

Aguirre et al. from MIT have also reported an electrostatic MEMS mirror based

endoscopic OCT probe [63]. In their paper, they demonstrated a gimbaled 2D MEMS

mirror design based on angular vertical comb actuators, which allows larger scan

angles compared to other vertical comb drives with the same dimensions. The mirror

has a circular aperture, whose diameter is 1 mm and the device footprint is 3 × 3 mm2,

as shown in Figure 1-8 A. Mechanical scan angles of ±6°have been achieved on both

axes at more than 100 V. They reported a side-view probe with an aluminum holder

housing the MEMS mirror. The probe has a diameter of about 5 mm. The optical

A B

C D

E F

28

components packaged inside the probe include a fiber, a GRIN lens and a small

achromatic lens (Figure 1-8 B). Incorporated in a spectral domain OCT, the probe scans

at 4 frames/sec over the range of 1.8 mm × 1.0 mm × 1.3 mm. 3D hamster cheek pouch

images have been obtained to demonstrate the capability of the probe.

Figure 1-8. Electrostatic MEMS OCT from MIT. A) SEMs of the MEMS mirror, B) design and packaging of the probe, and C) 3D image of hamster cheek pouch.

As is shown in Figure 1-7 B, the probe size is limited by the footprint of the MEMS

mirror. For electrostatic mirrors, much space of the device is taken by the comb drive

actuators. This results in a small ratio of the mirror aperture size to the device footprint

size, or small fill-factor for electrostatic mirrors, which may limit further miniaturization of

endoscopic probes. High resonant frequencies make them ideal for fast scanning

applications, but working at resonance can cause non-linear scan. High driving voltage

required may also be a concern for endoscopic applications.

Electromagnetic MEMS Endoscopes

To further increase the scanning range and lower the driving voltage,

electromagnetic mirrors are explored for endoscopic OCT applications. Electromagnetic

actuation is based on Lorentz force; larger driving force can be realized with lower

driving voltage. By controlling the current flowing direction, both repulsive and attractive

driving force can be realized [64]. The magnetic field for electromagnetic actuation is

normally generated by permalloy [64]-[66] or active electric coils [64], [67]-[68]. A 1D

A B C

1mm

Inner torsion beam

AVC actuators

Outer torsion beam Gimbal

Y X

Control SMF FC lens Mirror

AL package

5 mm

29

micromirror with over 60° deflection angles has been reported by Miller et al. [64]. A 2D

micromirror with a mirror plate as large as 3.5 mm × 3.5 mm has been demonstrated.

Working at its resonance at 2 kHz, with 20 mA driving current, the mirror scan angle

could reach ±1.51°, and the movable frame can scan ±5.71° [69]. Yang et al. proposed

a coil-less design, in which the optical scan angle can reach 20° with a 2 mm × 2 mm

mirror plate when the mirror was operated at an input power of 9 mW [70].

Kim et al. demonstrated a 2D electromagnetic MEMS mirror based endoscopic

OCT probe [71]. A gimbaled 2D mirror design was employed. A permanent magnet was

glued to the backside of the mirror plate, and wire-wound coils were placed inside the

probe body for each scan direction. The mirror plate was 0.6 mm × 0.8 mm and the

device footprint was 2.4 mm × 2.9 mm, as shown in Figure 1-9. About ±30° optical scan

angle was obtained with ±1.2 V and ±4 V driving voltages for the inner and outer axis,

corresponding to 50 mA and 100 mA current respectively. A probe with a 2.8 mm

diameter and 12 mm length has been demonstrated with SD-OCT. 3D images of finger

tips were obtained at 18.5 frames/sec. ±2.8 V and ±0.8 V voltages were applied on the

inner and outer axis, covering 1.5 mm × 1 mm lateral scan range, and consuming 150

mW power in total.

Watanabe et al. have recently demonstrated another electromagnetic MEMS OCT

probe [72]. The fabricated mirror module is shown in Figure 1-10. The mirror plate size

is 1.8 × 1.8 mm2, and the device chip is as large as 10 × 10 × 0.2 mm3 due to the large

electrical coil required. The entire device was placed on a 15 mm × 15 mm × 1 mm

PCB, which was fixed on the holder with a magnet inside. The MEMS mirror was used

in a Fourier domain OCT and 3D human finger images were obtained.

30

Figure 1-9. Electromagnetic MEMS OCT reported by Kim et al. [71]. A) SEMs of the MEMS mirror, B) design and packaging of the probe, and C) 3D image of fingertip.

Figure 1-10. Electromagnetic MEMS OCT from Yamagata Research Institute of Technology, Japan. [72] A) MEMS mirror module, B) MEMS probe in OCT system, and C) 3D image of human fingertip.

A C

B

A B

C

F E

D

Inner axis

Outer axis

Slow coil

Fast coil pair

MEMS

Fiber

Grin Grin lens Optical fiber Fold mirror

MEMS scanner

MEMS scanner

6 (typ.)

Torsion beams

Printed board (t 1 mm)

Y scan coils Au mirror

Flex lead

10

Magnet and holder X scan coils

X

Y 15

Epidermis

Dermis

31

One of the drawbacks of electromagnetic mirrors for endoscopic application is its

high power consumption. The other drawbacks lie in the fact that external magnets are

required for actuation, which not only greatly complicates the packaging process, but

also constraints the further miniaturization of the probes. Electromagnetic interference is

also a concern.

Piezoelectric MEMS Endoscopes

Piezoelectric actuation has also attracted some attention. It takes advantage of the

piezoelectric effect, and realizes bending motion by applying electric field across a

piezoelectric material such as lead zirconate titanate (PZT). The advantages of

piezoelectric actuation include fast response, large bandwidth and low power

consumption. Piezoelectric actuators are usually composed of metal/PZT/metal

sandwich [73]-[76] or double layered PZT materials [77]. Scanning angles as large as

40° have been reported, using voltages up to 13V [78].

A piezoelectric MEMS based OCT probe has also been reported [79].

Piezoelectric MEMS mirrors with aperture sizes of 600 µm × 840 µm and 840 µm ×

1600 µm have been fabricated. Mechanical scan angles up to ±7° and resonant

frequency up to 1 kHz were measured for the mirrors. A prove-of-concept probe design

is shown in Figure 1-11 B. The 600 µm × 840 µm mirror was used in a FD-OCT system

to demonstrate its imaging capability, and a 2D image (Figure 1-11 C) of an IR card was

obtained.

However, the mirrors are only one dimensional and two mirrors scanning at

orthogonal directions are required for 3D imaging. In addition, the large initial tilt angle

will complicate optical alignment and probe packaging. Furthermore, for piezoelectric

32

MEMS mirrors to be used for in vivo OCT imaging applications, one also needs to

overcome the charge leakage problems and hysteresis effect.

Figure 1-11. Piezoelectric MEMS OCT reported by Gilchrist et al. [79] A) SEM of MEMS mirror, B) 3D MEMS probe design, and C) 2D image of IR card.

Electrothermal MEMS Endoscopes

Eletrothermal actuation is studied to further increase the scanning range at low

driving voltage. Electrothermal actuation can be realized by using bimorph beams. A

bimorph beam is formed by two layers of materials with different thermal expansion

coefficients. The bending motion of the beam is induced by the expansion difference of

the two materials in response to temperature change. The actuation force typically is

larger than that of electrostatic or electromagnetic actuation. Electrothermal

micromirrors also have almost linear response between the scan angle and applied

voltage, and other advantages, such as simple structure design, and easy fabrication.

One of the most important features of electrothermal mirrors is their high fill factor. For

the same mirror aperture size, the device can be smaller, which is crucial for making

smaller endoscopic probes. Different materials have been explored for fabricating

bimorph beams [80]-[87], including silicon (Si), various metals, such as aluminum (Al),

polymers, and dielectrics. The most popular choice is Al and SiO2. It was first

demonstrated by Bühler et al. in 1995 [81]. Then Jain et al. demonstrated 2D

B C A

33

micromirrors based on Al/SiO2 bimorph beams and polysilicon as the heating material

with rotation angle up to 40°on a 1 × 1 mm2 mirror plate at 15 V [88].

Electrothermal actuation was the first of the four actuation mechanisms to be used

in endoscopic OCT imaging [45], [89]. Xie et al. reported a 1D SCS MEMS mirror. The

mirror aperture size was 1 mm × 1 mm, and 17° rotation angle and 165 Hz resonance

frequency were measured. A forward-looking probe was demonstrated and applied for

in vivo imaging of porcine urinary bladders. 5 frames/sec imaging speed was achieved

with 2.9 mm × 2.8 mm imaging area. Yet, due to the mesh actuator design, a 10° scan

discontinuity was observed. Later in 2003, the actuator design was improved by

replacing it with a paralleled actuator beam structure, resulting in continuous angular

scanning with the scan angle increased to 37° [90]. 2D ex vivo imaging of a rabbit

bladder, as shown in Figure 1-12 E, was obtained to demonstrate the probe capability.

Figure 1-12. Electrothermal MEMS OCT reported by Pan et al. A) SEM of 1D MEMS mirror, B) Meshed actuator structure, C) parallel beam structure, D)OCT system with forward looking endoscopic probe design, and E) 2D OCT image of rabbit bladder.

A

B

D

C E

34

Figure 1-13. Electrothermal MEMS OCT reported by Xu et al. A) two-axis electrothermal MEMS mirror, B) probe design and assembly demonstration, and C) 2D and 3D OCT image of an IR card.

More recently, Xu et al. also reported an electrothermal actuation based MEMS

OCT probe [91]-[94]. They have demonstrated a series of MEMS mirrors with both

straight and curled shaped electrothermal bimorph actuators formed by Aluminum and

Silicon. The largest mechanical deflection reported was 17° at an operation voltage of

~1.3V [91]. The device has a 500 µm diameter mirror aperture on top of a 1.5 mm × 1.5

mm chip. The mirror showed good linearity between driving voltage and scan angle

after the initial critical voltage, and the 3 dB cutoff frequency was 46 Hz. The probe

assembly was based on silicon optical bench (SiOB) methodology for self-alignment of

the optical components, and the electrical connection from the MEMS mirror to the

copper wires on the substrate was made using solder balls. The probe was then

inserted into a transparent housing. The diameter of the probe was less than 4 mm, and

the rigid length was about 25 mm. The probe was used in a swept source OCT system

at a frame rate of 21.5 frames/sec for 3D imaging; and the imaging results of an IR card

were shown in Figure 1-13C.

An improved design based on Xu's report was later reported by Mu et al. [95]. The

mirror scanning angle was 11° at less than 2 V driving voltage, and the mirror plate was

increased to 1 mm in diameter. The probe diameter shrank to 3mm, and a similar SiOB

A B C

35

assembly process was adopted. 2D images of mouse muscle and skin were

successfully demonstrated.

Of the four actuation mechanisms, electrostatic actuation has fast response and

lowest power consumption, but it requires large driving voltage, which may not be safe

for imaging human internal organs. Electromagnetic actuation can realize large scan

angle, at low driving voltage, but the permanent magnet required poses difficulties for

further probe miniaturization. Piezoelectric MEMS mirrors can also be fast and consume

low power, but it must overcome the large initial tilting issues, the hysteresis effect and

charge leakage problem for OCT imaging applications. Electrothermal actuation can

scan large angles at low driving voltages, and it also has the largest fill-factor compared

to all three other types of MEMS mirrors. Thermal response is relatively slow, but it is

capable to achieve real time imaging. Overall, electrothermal MEMS mirrors are the

better choice for endoscopic OCT scanners.

Research Goals

This research explores the applications of OCT technique using MEMS scanners.

One major task is to develop endoscopic OCT probes based on electrothermal MEMS

mirrors for tissue imaging, and prove its ability for early cancer detection. The other one

is to employ OCT for localized brain tissue refractive index (RI) measurement. This

study could help to correct optical image distortion caused by heterogeneous tissue

structures or distorted tissue image caused by disease or surgical loads. Another

research goal is to realize functional OCT, including polarization sensitive OCT and

Doppler OCT to improve image contrast for birefrigent tissues and blood vessels. Lastly,

to develop MEMS based endoscopic probes that are suitable for both OCT imaging and

photoacoustic imaging; photoacoustic imaging can achieve higher imaging depth at a

36

cost of lower resolution, by combining the advantages of the two imaging modalities,

one imaging session would yield more information for diagnosis.

Thesis Outline

There are seven chapters in this thesis. Chapter 1 introduces the current imaging

techniques and their limitations in the application of early cancer detection; a solution

based on MEMS technology, specifically eletrothermal MEMS mirrors as the scanning

engine has been proposed. Background and literature reviews are also given in this

chapter. Chapter 2 focuses on introducing OCT principles, including both time domain

and frequency domain systems. Functional OCT: Doppler OCT and polarization

sensitive OCT will also be introduced. Chapter 3 first introduces electrothermal

actuation mechanism, and then goes on to introduce four generations of MEMS mirrors

developed at BML, including two generations employed in this work. Chapter 4 presents

endoscopic MEMS OCT imaging. Four generations of probe designs and their assembly

processes will be introduced; imaging results obtained with assembled probes will be

presented to demonstrate the probe capabilities. And the optical qualities of the probes

will be analyzed. Chapter 5 discusses OCT's application in measuring localized

refractive index in brain tissue slices using MEMS mirror for lateral scans. RIs measured

in five different regions and how RIs change in grey matter and white matter under

compression will be reported. Chapter 6 covers functional OCT. PS OCT and Doppler

OCT results will be presented in this chapter. Chapter 7 discusses combining OCT with

photoacoustic imaging. Preliminary results will be presented. And lastly, in Chapter 8, a

summary will be given, and future work will be listed.

37

CHAPTER 2 OPTICAL COHERENCE TOMOGRAPHY

OCT was first introduced in the early 90s at MIT [2]. Due to its high resolution, and

relatively large penetration depth, OCT has become the ideal imaging modality for non-

invasive early cancer detection. Over the years, OCT has evolved from time domain to

frequency domain, resulting in much faster systems; Functional OCT has also been

developed. Doppler OCT helps to image the cardiovascular system with better contrast;

while PS OCT targets regions with birefringent properties. In this chapter, the principle

of low coherence interferometry will be introduced first. Then time domain and

frequency domain OCT implementations and their difference will be discussed. For time

domain OCT, different depth scanning mechanisms will be reviewed. Finally, functional

OCT principles will be presented.

Low Coherence Interferometry

The core of OCT system is a low coherence interferometer. The most popular

configuration of OCT system is a Michelson interferometer, which is shown in Figure 2-1

A. Michelson interferometer is a kind of division-of-amplitude interferometer. The input

light is divided into two arms at the same point in space by a beam splitter and travels

through different paths from that point on. Reflective mirrors are placed at each end of

the arms, and input signals are reflected by the mirrors at the ends and travels back to

recombine at the beam splitter. When two beams of light have the same frequency,

same vibration direction, and a certain phase difference, interference pattern can be

observed at the photodetector. Simple math can help prove this statement. The electric

fields of two beams of input light can be expressed as

)cos( 111111 trkAE

(2-1)

38

)cos( 222222 trkAE

(2-2)

In the equations, A

is the field amplitude vector;

2k is wave number; is the

frequency; and represent the initial phase of the electrical field. When the two beams

combine at one spatial point, the intensity of the combined signal can be expressed as

cos22

2)()(

21212121

12212122112121

AAIIEEII

IIIEEEEEEEEEEI

(2-3)

)()()( 21212211 trkrk is the phase difference of the two signals. 1I

and 2I are the individual intensities of the two input signals, and 12I is the interference

term. One prerequisite for the interference term to exist is that the two beams have

parallel amplitude portion, or else the two amplitude vectors are perpendicular to each

other, which will result in 0cos2121 AAAA

. The other requirements for obtaining

a stable interference pattern are: first, the phase difference should not change with time,

meaning 21 ; and second, the initial phase difference 21 should be fixed. Since

the two beams from Michelson interferometer are from the same light source, these

requirements can be satisfied; as a result, stable interference pattern can be observed

at the photodetector.

When the input light source is of single wavelength , we could obtain

221 kk , then the question could be simplified if we consider using a 50/50 beam

splitter, the two beams will have the same amplitude 21 AA , and the same

intensity 021 III . As a result, the interference signal can be expressed as

39

)4

cos1(2 0 rII

. r is the optical path difference between the two arms, and the

term 4 is obtained here, because the light goes through the optical paths two times.

Figure 2-1. Michelson interferometry. A) Michelson Interferometer setup, B) Interference

pattern of wavelength .

The interference pattern of monochrome light source is shown in Figure 2-1B. It is

a cosine wave with a period of 2

. Theoretically, the interference length is unlimited. But

in reality, all light sources have their own bandwidths, expressed as . Different

wavelengths will have different interference periods, and each interference pattern will

add up to result in a final interference pattern with a limited interference length, as

shown in Figure 2-2.

Figure 2-2. Gaussian distribution broadband light source interference pattern.

For a light source with a Gaussian spectrum, the intensity of the interference

pattern also has Gaussian distribution. In this case, the coherence length can be

expressed in the following form:

22

44.02ln2

cl . Only when the optical path

B A Photodetector

Beam splitter

0

x

I

40

lengths of the two arms are matched to within the coherence length, can interference

pattern be generated. It is clear from the equation that the coherence length is inversely

proportional to the bandwidth of the light source. By employing broadband light source,

the coherence length can be effective reduced. And this is the principle based on which,

ultrahigh axial resolution of OCT system could be realized.

Time Domain OCT System

In a time domain OCT system, depending on the functions of the arms, one of

them is chosen to be the reference arm, while the other one is the sample arm. A

movable reflective mirror, whose position can be controlled precisely, is placed at the

end of the reference arm; while a sample of interest is placed at the end of the sample

arm. The biological samples can be deemed as a series of tissue layers, the

backscattered light from each layer contains information of that specific depth. The

reference mirror scans through each depth, and only the signals from the depths that

match with the reference arm optical path length to within the coherence length will be

collected, while the other signals are rejected by coherence gating.

Figure 2-3. OCT sample signal.

Beam splitter

Photodetector

Reference arm

Sample

41

Superluminescent Diodes (SLDs) are usually employed by time domain OCT,

because they can offer high brightness, and low coherence length [96]. Commercially

available SLDs at the near infrared range has bandwidth from 45 nm to 110 nm, with

emitting power over 10 mW, resulting in axial resolution as small as 7 µm in air [97].

Efforts have also been made to combine multiple SLDs to form a light source with

broader bandwidth [98]. One SLD light source combining two SLDs has already been

commercialized by Thorlab, the bandwidth reached 200 nm, with center wavelength of

1300 nm, axial resolution as high as about 3.7 µm in air can be obtained [99].

Depth scanning, also called A-Scan, is performed by moving the reference mirror

back and forth over a desired imaging range. In early OCT systems, a linearly

translating galvanometer is used for depth scan [2]. This method is simple and

straightforward, but its low speed prevented it from being used in real time imaging

applications. To increase the scanning speed, Tearney et al. proposed a fiber stretching

based optical delay line, which is realized by employing piezoelectric actuation [100].

Although high scanning speed is obtained, this method has some disadvantages. First,

it requires high driving voltage; second, non-linear translation occurs due to hysteresis

effect; third, there is polarization mode dispersion, which will decrease the image

resolution. In these two methods, the movement of the reference mirrors will induce a

carrier frequency for the interference signal. This is desirable for a better signal to noise

ratio, since the transmission of the signal can avoid the base band where flicker noise

dominants. Later, a grating based rapid scanning optical delay line (RSOD) is

introduced [101]-[108]. This method is based on the technique of femtosecond Fourier

transform pulse shaping. The input light is optically Fourier transformed and modulated

42

in frequency domain. Early attempt for pulse shaping employed a symmetric

configuration; phase modulation is realized by a liquid-crystal phase modulator, shown

in Figure 2-4 A [103]. Later, by employing a double pass mirror, a more compact RSOD

design is introduced [106], [107], shown in Figure 2-4 B [107]. Linear phase modulation

is realized by rotating a scanning mirror at very fast speed. Scanning rate of 6 m/s and a

repetition rate of 2 kHz were realized. This method can be used to control group delay

and phase delay separately, and also is able to achieve high scanning speed, high

repetition linear scanning. This method could be used for systems that requires no

dispersion. On the other hand, by changing the position of the grating, certain

dispersion could be introduced into the system, which could be used to cancel out the

dispersion in sample arm.

Figure 2-4. RSOD configurations. A) symmetric pulse shaping configuration, B) folded RSOD design employing double pass mirror [106].

A

B

43

Figure 2-5. Examples of applications of time domain OCT system. A) OCT imaging of rabbit esophagus in vivo. (m: mucosa, sm: submucosa, im: the inner muscular layer, om: the outer muscular layer, s: serosa, a: adipose and vascular supportive tissues [104], B) OCT imaging of rabbit bladder. (U: urothelium, LP: lamina propria, M: muscularis, F: attached fatty tissue)[106].

Frequency Domain OCT

To further increase depth scanning speed, Fourier domain OCT (FD-OCT) system

was introduced [109]. In FD-OCT, no mechanical depth scan is required, instead, the

spectrum of the interference signal is obtained, and the depth information can be

calculated via inverse Fourier transform. FD-OCT can be implemented in two ways.

One way is to use a diffraction grating to generate the spectrum of the interference

signal spatially, and the signal is detected by a spectrometer. Using this implementation,

the depth scan is finished during the time for one exposure of spectrum. The other

method uses an ultra-fast sweeping light source to temporally disperse the signal; one

photo detector is needed to capture the signal. Depth scan of this implementation is

done during one wavelength sweeping period. Swept laser sources that have scanning

frequency of up to 20 KHz over a tuning range of 120 nm full width has been reported

[110]. Applied in OCT system, one depth scan can be done in 250 µs, greatly increased

the imaging speed. FD-OCT has been reported to have good signal to noise ratios;

sensitivity of around 90 dB have been achieved [111], [112]. Many research groups

A B

44

have successfully demonstrated FD-OCT, images of human retina, fingertip, etc. have

been obtained [112]-[114].

Figure 2-6. Example application of FD-OCT system: tomographic scan of the entire left circumflex coronary artery obtained with frequency domain OCT system (C7xR, LightLab, Watford, USA). A) longitudinal view of 54 mm of the artery at the pullback rate of 20 mm/s, B-E) multiple cross sectional images selectively obtained at different levels of the native left circumflex artery [116].

Principle of Frequency Domain OCT

FD-OCT is based on spectral interferometry technique [117]-[119]. It employs the

same system configuration as TD-OCT; the difference is that the reference arm is fixed

in this case, with an optical path length difference with the sample arm.

Consider the sample arm only, when incident with a plane monochromatic wave,

whose electric field can be expressed as )exp( tirkiAE iii

, based on the Fourier

diffraction theorem, in far field, the electric field of the detected back-scattered light can

be related to the Fourier transform of the scattering potential of the object )(rF

:

Aa

B C D E

45

)(~

)exp(.

)()exp(.)exp()()exp(.)( 3

KFtirkiC

rFFTtirkiCrdrKirFtirkiCkE

s

ssss

(2-4)

In this equation, .C is a constant, sk

is the wave vectors of the scattered light,

is kkK

is the scattering factor. When assuming the scattering potential F

is constant

over x and y direction, the equation can be simplified in the z direction, a new constant

.1C could be used to include the integral over x and y directions, while D is the

distance from the sample surface to the photodetector

)()exp(.1)exp()( zFFTtiDikCtiDikkAkE ssssss (2-5)

)(zFFTkA ss (2-6)

As is shown here, the scattering potential can be obtained by inverse Fourier

transform of amplitude of the electrical field. Yet, the amplitude of the signal cannot be

directly detected; instead, the power of the back scattered light can be measured

through photodetector. The inverse Fourier transform of the interference power results

in the autocorrelation of the scattering potential.

)(][)]([ 211 zFACFkAFTkIFT sssS

(2-7)

In an OCT system, when the reference mirror is placed in front of the sample with

a distance L , and the sample thickness is T, shown in Figure 2-7. The strong reflection

at the reference mirror can be expressed as rzzR , where R stands for the

reflectivity of the reference mirror. The scattering potential from an OCT system, is

rS zzRzFzF )()( (2-8)

46

The inverse Fourier transform of the power of interference signal yields the auto-

correlation of )(zF , the system scattering potential.

)()()()()( * zRzzRFzzRFzACFzACF rSrSFsF (2-9)

Figure 2-7. FD-OCT signal detection schematic.

)(zACFF is consist of four terms, the third term of the above equation is the true

reconstruction of the object structure, centered at rzz . In order for the

)( zzF rS term to not overlap with the other terms, the distance of from the reference

mirror to the sample surface L should be greater than the sample thickness T. As a

result, depth information of the sample can be obtained through inverse Fourier

transform of the power signal.

Configurations of FD-OCT

In order to perform Fourier transform, the relationship between signal intensity I

and wave number k should be obtained. As was mentioned in the beginning of section:

Principle of Frequency Domain OCT, two frequency domain OCT configurations have

been proposed for this purpose.

0

T L

Reference

Mirror

Z=Zr Beam Splitter

Photodetector

Light Source

Sample

47

One FD OCT configuration is shown in Figure 2-8 A. In this configuration, a

broadband light source is employed as in the time domain setup. After the reference

and sample signals recombine at the beam splitter, the signal is incident on a dispersive

grating. The signals of different frequencies are then spatially separated, and the signal

detection is realized by using a spectrometer.

Figure 2-8. FD-OCT configurations. A) spectral domain OCT, B). SS-OCT.

The other configuration is shown in Figure 2-8 B. This setup employs a frequency

sweeping light source, which sweeps the signal frequency over a broad band.

Therefore, this configuration is usually referred to as swept source OCT (SS-OCT). Only

one photodetector is needed for this configuration. The wavelength of input light change

linearly with time and the interference signal is collected during one sweeping period.

Broadband source Beam splitter

Reference mirror

CCD array detector

Lateral scanning

PC DAQ

Dispersive

Grating

A

Sweeping source Beam splitter

Reference mirror

Photo detector

Lateral scanning DAQ PC

B

48

SNR Advantage of FD-OCT

Without having to mechanically scan the reference arm, the speed advantage of

FD-OCT is obvious; the other advantage of FD-OCT over TD-OCT is better signal to

noise ratio (SNR) [120], [121]. In time domain system, the interference part of the signal

has electrical power defined as ][/2 2222 AEPPeS samplerefsignal , in which is the

quantum efficiency, e is the electron charge, and /hcE is the photon energy. This

reference and sample arm signal power can be expressed by their power spectral

density.

dSP samplerefsampleref )(,, (2-10)

An assumption can be made here: the reference spectral density )(refS is the

same as the light source spectral density )(S ; while the sample spectral

density )(sampleS is an attenuation of )(S , which can be noted as )(S . As a result,

the signal power can be written as

2222 /])([2 EdSeSsignal

(2-11)

There are three main noise sources in the OCT system, they are thermal noise,

shot noise and relative intensity noise. The noise spectrum is expressed in the following

equation.

]/[)(224

)( 22

2

HzAE

eP

E

Pe

R

kTfN coh

refref

fb

noise

(2-12)

In this equation, k is Boltzmann’s constant, T is the temperature in degrees

Kelvin, fbR is the value of the transimpedance amplifier feedback resistor, and coh is the

49

coherence time of the source, which is

cccoh

22

44.02ln2

. Under shot noise

limited condition, the second term of the noise is dominant, and the SNR in time domain

can be expressed as in equation (2-13) , in which and 0 are the source spectral

bandwidth and center frequency. BW is the signal passband, which is normally two

times the carrier frequency spectrum.

dSBWE

dSSNR

)(

])([

2/

2/

22/

2/

1 (2-13)

When this signal is spatially separated onto two detectors, each detector obtains

half the original bandwidth, the signal power is the coherent summation of the signal

over the spectrometer, while the noise power is the incoherent summation of the two.

The new SNR is

12/

0

0

2/

22/

0

0

2/

2 2])(5.0)(5.0[

])()([SNR

dSBWdSBWE

dSdSSNR

v

(2-14)

As we can see here, SNR increased by a factor of 2. The above analysis can be

extended to N detectors, which is the case for FD-OCT. The exact value of SNR

increased depends on the source's spectral density and spectral bandwidth per

detector.

Functional OCT

Another branch of OCT system is functional OCT. We focus on two types of

functional OCTs: Doppler OCT and polarization sensitive OCT (PS OCT). They extend

OCT's ability from obtaining micro-scale intensity images of subsurface structures to

imaging velocity or birefringence.

50

Doppler OCT

Doppler effect describes the phenomenon that when there is relative movement

between the source and the observer, the frequency received by the observer is

different from the source frequency, and the change in frequency is related to the

velocities of the source and the observer. In standard OCT imaging, the signal collected

by the photodetector is modulated by a carrier frequency. According to Doppler effect, if

the imaged sample has a relative velocity to the incident light, the detected carrier

frequency would change, and by detecting the changed carrier frequency, velocity of the

sample could be obtained. This is the principle of Doppler OCT. Doppler OCT is

normally used to identify blood vessels. As is shown in Figure 2-9, when the target

blood vessel is place at ( 90 ) to the incident beam with blood flow velocity of sV , a

relative velocity of cossV could be generated, and blood vessels could be detected.

Figure 2-9. Doppler OCT sample beam.

One way to calculate the Doppler shift frequency is based on short time Fourier

transform (STFT) algorithm [125]. Interference signal intensity received at the

photodetector is a function of time. Different time corresponds to different depth

through reference mirror scanning velocity, depth rvZ . Interference signal

))(2cos()()( tfftAtI Dr . The spectrogram of a signal )(tS can be estimated by

51

computing the squared magnitude of the STFT of the signal )(tS , as shown

here: 2

),(ˆ),(ˆ fISTFTfImspectrogra ii . i represents the i'th time window ),( ii .

Each time window has a pixel size of vl . The spectrogram is an estimation of the

power spectrum of the time domain signal. And is an estimation of the power

intensity at frequency during the i'th time window. For structural OCT image, intensity

is obtained by calculating the spectrogram at carrier frequency rf , and the number is

mapped back to its depth location through its time window. In order to obtain Doppler

images, real carrier frequency is needed for each time window. And Drc fff is

obtained by calculating the centroid of the measured power spectrum.

m

miODTmiODT

m

mc ffff ),(ˆ/),(ˆ

(2-15)

This STFT method is effective. Yet its limitation lies in the fact that the minimum

detectable Doppler frequency shift is inversely proportional to pixel window time [126],

[127]. This introduces tradeoff between detection sensitivity and imaging speed, as well

as spatial resolution.

To overcome this problem, a phase resolved signal processing method is

proposed [128]. As shown in Figure 2-10, this algorithm compares the phase change in

consecutive A-lines to compute the carrier frequency: Tf KKD /1 . Hilbert

transform is used to obtain phase information. The detected interference signal is

)cos()( ttAt , Hilbert transform can be applied to obtain the complex signal as

shown in equation (2-16).

)()()sin()()cos()(

)()()(

~ tietAttiAttAdt

Pi

tt

(2-16)

52

Figure 2-10. Consecutive A-Line scans in Doppler OCT

Doppler frequency shift is determined by the average of n numbers of A-lines

frequency shifts, each one is expressed in the following equation, j is the number of A-

line, and T is the time between each A-line.

))((

1

))()((

1

*

1 )()()()()(~

)(~ 1 ti

jj

tti

jjjjjjj etAtAetAtAtt

(2-17)

)))(

~)(

~Re(

))(~

)(~

Im((tan

2

1

2

)()(

*

1

*

11

tt

tt

TT

ttf

jj

jjj

j

(2-18)

Polarization Sensitive OCT

Certain tissues such as retina, meniscus, tendons etc. have birefringence

property. When such types of tissues are injured, this property will be damaged or lost.

By detecting the polarization states of reflected light, birefringence information of the

sample can be obtained. PSOCT has been reported to be used for measuring the

sample reflectivity, phase retardation, Stokes parameters, Muller and Jones matrix,

birefringent axis orientation and diattenuation [129]-[138].

Most of early works are based on the setup proposed by Hee et al. [129], system

schematic is shown in Figure 2-11. Input light from the light source is linearly polarized

by going through a polarizer. Horizontally polarized light can be described

Axial scanning

Lateral scanning

53

as

0

10EE using Jones Matrix, in which tikAE exp)(00 . This linearly polarized

light is then split into reference arm and sample arm by a 50/50 beam splitter, written

as

0

1

2

100 EES and

0

1

2

100 EER separately. A quarter-wave plate, with its fast

axis 22.5° to the X-axis is inserted in the reference arm. Light goes through it twice as it

is reflected from the reference mirror, the slow axis has a phase delay of 180°

compared to the fast axis, and the output light from the reference arm is linearly

polarized with 45° to the X-axis, )2exp(1

1

2

101 rR kliEE

. rl is the travel distance from

the beam splitter to the reference mirror. The amplitude of X-axis and Y-axis are equal

from the reference arm. In the sample arm, a quarter-wave plate with its fast axis 45° to

the X-axis is used, circularly polarized light is incident upon the sample, the Jones

Matrix for the quarter wave plate is

1

1

2

1

i

iJW . Assuming the birefringent axis of the

sample has an angle to the X-axis, and the phase retardation between the fast and

slow axis is . The Jones Matrix of the sample can be expressed as

)(sin)2

exp()(cos)2

exp()2

sin()cos()sin(2

)2

sin()cos()sin(2)(sin)2

exp()(cos)2

exp(

22

22

iii

iii

J s

(2-19)

Using Jones Matrix, the light coming back from the sample can be expressed as

)2exp()()cos(

)sin()2exp(

2

1

)2exp()0

1(

2

1)()(

0

01

sS

sSWSSWS

kliEdRi

kliEJJdRJJdE

(2-20)

54

In this equation, dRS is the sample's reflectivity, sl is the sample path length,

which equals 0l , the distance from the beam splitter to the sample surface, plus d , the

depth into the sample; as shown in this equation dlls 0 .

Figure 2-11. Free space PSOCT system setup.

The two beams of light recombine at the beam splitter, and are divided by the

polarization beam splitter (PBS2) into two orthogonal polarization states. The

interference signals of each polarization state are detected by two photodetectors

separately. The horizontal and vertical signals at the photodetectors are shown in the

following equations.

0)2exp()sin()2exp()(2

)2exp(

2

1EkliidR

kliE sS

rH

(2-21)

0)2exp()cos()(2

)2exp(

2

1EklidR

kliE sS

rV

(2-22)

The current detected at the photodectors are shown in equation (2-23) and

equation (2-24). is the detector sensitivity, which is assumed to be the same for both

55

detectors; and l is the path lengths difference between reference and sample

arms sr lll .

8

)()22cos()sin()(22)(sin)(21

2

022 kAlkdRdREi SSHH

(2-23)

8

)()2cos()cos()(22)(cos)(21

2

022 kAlkdRdREi SSVV

(2-24)

The cross-correlation part of the signal is the interference signal,

22

)()22cos()sin()(

2

0 kAlkdRi SHac

(2-25)

22

)()2cos()cos()(

2

0 kAlkdRi SVac

(2-26)

Assuming the light source has a Gaussian power distribution, )(kS is the spectral

density, and is the FWHM,

2

02ln2l is the coherence length, integrating the

signal over the light spectrum gives

)22cos()sin(

])2ln2

(exp[)(22

)()22cos(

)sin()(22

0

20

0

lk

l

ldR

PdkkSlk

dRP

dkiI

S

SHH ac

(2-27)

)2cos()cos(

])2ln2

(exp[)(22

)()2cos(

)cos()(22

0

20

0

lk

l

ldR

PdkkSlk

dRP

dkiI

S

SVV ac

(2-28)

The signal at this point is demodulated at the carrier frequency to eliminate

the lk 02cos term,

56

)sin(2ln2

exp)(22

2

0

l

ldR

PI SH

(2-29)

)cos(2ln2

exp)(22

2

0

l

ldR

PI SV

(2-30)

Let

2

2ln2exp)()(

l

ldRdR SSr

, and by eliminating the constant term

22

0P, the signals can be expressed by the following equations.

)sin()(22

0

dRIP

S SrHH

(2-31)

)cos()(22

0

dRIP

S SrVV

(2-32)

The reflectivity of the tissue can then be calculated from22

)( VHSr SSdR , and

the phase retardation is

V

H

S

Sarctan .

As is shown in equation (2-27) and equation (2-28), the birefringence axis

orientation is only related to horizontal polarization. So to obtain , we need to obtain

the phase difference between vertical and horizontal polarization. This information can

be obtained by signal processing introduced in Section: Doppler OCT. Using Hilbert

transform, 22 0 lkH and lkV 02

can be obtained, and

2

180 VH

can

be calculated.

57

Time Domain OCT System in BML

A time domain OCT system is employed in BML. The schematic of the system is

shown in Figure 2-12. The employed broadband light source (DenseLight, DL-BX9-

CS3159A) has a minimum power of 15 mW, its center wavelength is 1310 nm, and its

FWHM (full width half maximum) is 75 nm, resulting in a 10 µm axial resolution in air.

The input light is divided into the reference arm and the sample arm by a beam splitter.

Depth scanning at the reference arm is realized by a rapid scanning optical delay line

(RSOD), the scanning range is 0 to 1.6mm depth. The RSOD employs a galvanometer

scanning at 1 kHz. We use a 500 kHz carrier frequency for the OCT signal, and it is

generated by the small misalignment between the galvanometer center axis and the

optical axis. MEMS based probes are used in the sample arm to perform 2D transverse

scanning, which will be introduced in detail in Chapter 3. The OCT signal is detected by

a balanced photodetector, acquired by a DAQ card (NI PCI-5122), and processed by a

computer. The frame rate of our system is 2.5 frames/sec, and the sensitivity is 74 dB.

Figure 2-12. OCT system schematic.

Here, a detailed description of the RSOD will be presented. The grating based

RSOD employs a blazed diffraction grating, an imaging lens, a scanning mirror and a

Light source 2×2 Beam

splitter Circulator

RSOD

Photo detector

Endoscopic probe DAQ PC

58

double pass mirror. The RSOD configuration is shown in Figure 2-13. This design is

based on optical Fourier transform. When diffraction grating is placed at the front focal

plane, the Fourier transform of the optical signal can be obtained at the imaging focal

plane. One of Fourier transforms' properties is that a linear phase ramp in the frequency

domain corresponds to a delay in the time domain. The phase change can be induced

by tilting the scanning mirror. In this case, the signal is modulated in its Fourier domain,

and then transformed back to time domain to realize depth scanning.

Figure 2-13. RSOD configuration

Figure 2-14. Ray trace of RSOD with no center misalignment. A) Green light representing center wavelength, B) Blue light representing an additional wavelength.

A B

59

The scanning depth can be calculated through optical ray trace. The broadband

incident light is dispersed by the blazed grating and its spectrum is imaged on the

scanning mirror. The initial position of the grating, lens and scanning mirror are parallel

to each other. The incident angle of the light beam to the grating is , the diffraction

angle is . The blazed grating equation is mp sinsin , p is the grating constant,

which equals to mm300

1 in our system, m is the diffraction order, normally 1m . Take

the green light as the center wavelength 0 for example, as shown in Figure 2-14 A. The

RSOD is designed to give 0 a 0 diffraction angle, which means 00 , so we obtain

0sin p . In our system, nm13100 , so

14.23arcsin 0

p

. When the scanning

mirror has a tilting angle of , its optical path lengths change compare to 0 tilting one

way can be expressed by the following equation

p

f

pfff 00 2

2sin2sin)2sin(

(2-33)

Due to the double pass design and the direction in the graph, the actual scanning

range is p

f04 . In our system, the scanning depth is designed to be 1.6 mm, so the

tilting angle of the galvanometer should be 97.0017.0 rad . Now consider

dispersion, when there is a second wavelength as shown in Figure 2-14 B. Its

scanning range can be expressed as

fp

f4

4 0 . We can obtain 0sinp ,

based on the grating equation. When is small, p

0 and its scanning range can

be expressed as

60

)2(4

44

0

00

p

f

pf

p

fZ scan

(2-34)

The phase change induced is

12

422 0

p

fzscan . Knowing

that

2 , can also be expressed as

12

42

0

p

f. At this point, we can see

that by tilting the scanning mirror, phase change induced is in linear proportion to light

frequency. Delay in time domain is expressed by group delay, which is defined

as

d

dg

.

So by employing the grating based RSOD, dispersion can be eliminated, and a

constant group delay linear to scanning angle is generated throughout the spectrum,

which is0

16

p

f . When there is a misalignment between the center of the scanning

mirror and its axis of rotation, a carrier frequency can be generated. For a source with a

symmetric spectral distribution, carrier frequency is defined by the temporal rate of

phase change at the center wavelength. When the misalignment is x , carrier frequency

in RSOD can be expressed as scan

carriert

xf

14

. This parameter is linear to

misalignment x . In our system, carrier frequency is designed to be 500 kHz, and the

galvanometer scanning frequency is 1 kHz, so the period of the galvanometer is 1 ms,

and mstscan 5.0 , the misalignment mmx 8.4 . When the system is perfectly aligned, that

is 0x , there will be no carrier frequency, no matter what the scanning angle is.

Therefore, the phase delay and group delay can be separated, and controlled

individually. Spectrum of the carrier frequency induced by this method can be obtained

61

by calculating the derivative of carrier frequency with respect to wavelength, as in

2

4

scan

carrier

t

x

d

dfBandwidth . The bandwidth is about 28 kHz in our system. In

signal processing, filter bandwidth is normally 2 to 2.5 times of this bandwidth to avoid

the loss of signal, while keeping the system sensitivity.

Summary

In this chapter, the principle of OCT system has been introduced. OCT imaging

could be conducted in both time domain and frequency domain. Different depth

scanning mechanisms for time domain OCT system have been reviewed. Two types of

frequency domain OCT system configurations have been introduced. The differences

between time domain and frequency domain OCT have also been discussed. Two types

of functional OCT have been presented in this chapter. The principle of Doppler OCT

was introduced along with two different algorithms for extracting Doppler information.

For PS OCT, its principle and background were presented. Finally, the configuration

and performance of the time domain OCT system employed in this study was

introduced in detail.

62

CHAPTER 3 ELECTROTHERMALLY ACTUATED MEMS MIRRORS

A series of MEMS mirrors based on electrothermal actuation mechanism have

been developed in BML aiming to be applied in endoscopic OCT imaging applications.

This chapter first introduces the principle of electrothermal actuation, and then goes on

to introduce four generations of BML electrothermal MEMS mirrors; their designs and

characteristics will be presented.

Electrothermal Actuation

As mentioned briefly in Chapter 1, electrothermal actuation is based on bending

motion of bimorph beams. Bimorph beams are cantilever beams that are composed of

two different kinds of materials. Due to the difference in thermal expansion coefficients

in the two materials, when heated, the bimorph beam will bend. And that is the basic

principle of electrothermal actuation. Joule heating effect describes the phenomenon

that heat is generated when current goes through a resister; it is normally the source to

provide heat to bend the bimorph beams.

Figure 3-1 depicts the basic geometrical relationships between the substrate, the

bimorph beams, and the rigid frame of one actuator. The materials chosen for bimorph

beams are Al and SiO2, and Pt is chosen as the heating material. After the beam

structure is released, the residue stress will cause the bimorph beam to curl up, the

initial tilting angle can be calculated by

L

, as shown in Figure 3-1, l is the length of

the bimorph, is the tilting angle, and is the radius of curvature of the beam. When

applied with voltage, the actuator will be heated up due to Joule heating effect, Al has a

bigger thermal expansion coefficient than SiO2, which will cause the bimorph beam to

63

bend downwards. The scanning angle is proportional to the power, so it can be

controlled through controlling driving voltage for the actuator.

Figure 3-1. Eletrothermal bimorph structure.

1st and 2nd Generations Electrothermal MEMS Mirrors

Several actuator or MEMS mirror designs based on this structure have been

proposed, including both 1D and 2D mirrors. The MEMS mirror shown in Figure 3-2 A

has 1D scanning capability, it employed a series of bimorph beams, and the solid frame

is sputtered with Al as the mirror surface. To add a second scanning dimension, another

set of bimorph beams are added to the frame in the direction that is perpendicular to the

first set, as is shown in Figure 3-2 B [139]. 2D scanning can be realized by controlling

the 2 sets of bimorphs separately. For this design, the inner axis can scan up to 40° at

15 Vdc, and the outer axis can scan 25° at 17 Vdc. These mirrors belong to the 1st

generation mirror design.

Al Si SiO2 Pt

on actuation

64

Figure 3-2. 1st generation MEMS mirrors. A) 1D MEMS mirror, B) 2D MEMS mirror.

Despite the large scanning angle achieved by the 1st generation mirrors, the

drawback is obvious too. The center of the mirror plate shifts as the mirror rotates,

making optical alignment difficult and can cause artifacts in the imaging results. To

solve this problem, the 2nd generation mirror design was proposed by adding a set of

compensating bimorph beams to the frame on the same axis as the mirror anchor, as

shown in Figure 3-3. By matching the driving voltages for the two sets of bimorph

beams, mirror plate center shift can be eliminated. However, only piston motion can be

realized, therefore, it is not suitable for OCT application, which requires 2D scanning

capability. To add another scan dimension in order to realize tip-tilt motion, another set

of bimorph can be added on the perpendicular axis, similar to the previous design. Yet

the resulting structure is too complex, and the fill factor is very small, only around 5%.

The other disadvantage that makes this design not practical is that due to the variation

in fabrication process, the resistance values of actuators vary; each mirror has its own

driving voltage ratio which needs to be measured individually through experiment. A

more practical 2D MEMS mirror design suitable for OCT is therefore desirable.

A B

65

Figure 3-3. 2nd generation: lateral shift free (LSF) MEMS mirror design. A) LSF bimorph actuator design, B) top view model of 1D LSF MEMS mirror, C) SEM of 1D LSF mirror, and D) SEM of 2D LSF MEMS mirror.

LSF-LVD (3rd) and FDSB (4th) Electrothermal MEMS Mirrors

The 3rd and 4th generations of electrothermal MEMS mirror designs were later

proposed for endoscopic OCT imaging applications. They are the Lateral-Shift-Free

(LSF) Large-Veridical-Displacement (LVD) design and the folded-dual-S-shaped

bimorph (FDSB) design. The two designs and their characteristics will be reviewed in

this section.

LSF-LVD MEMS Mirror

LSF-LVD design is proposed to solve the design problems mentioned in the last

section. It is named Lateral-Shift-Free (LSF) Large-Veridical-Displacement (LVD)

because in this design, the mirror center does not shift during scanning, and large

vertical displacement of the mirror plate can be realized [140]. Each LSF-LVD actuator

is composed of three Al/SiO2 bimorph beams and two rigid silicon frames. Pt is

A B

D C

66

embedded as the heater. Each actuator is connected to a bonding pad on the front

surface of the mirror substrate. And they all share a common ground, which is

connected to the middle pad. Electrical connections to the actuators can be made

through the pads. The folded design effectively shrunk the area size of the actuators,

making high fill factor possible. In order for the structure to compensate the lateral shift

and tilting during vertical displacement, the dimensions of the five components of each

actuator need to meet the requirements established in equation (3-1). As is shown in

Figure 3-4 A, L1, L2 are the lengths of the two frames, and 1l , 2l , 3l are the lengths of

the three bimorph beams.

21

2312

1

LL

lll

(3-1)

Figure 3-4. LSF-LVD eletrothermal actuator and mirror design. A) side view of actuator and relationship between each element, B) top view of the MEMS mirror.

The mirror plate is held by four identical actuators. The size of the mirror plate is 1

mm × 1 mm, and the substrate is 2 mm × 2 mm, resulting in a high fill factor of 25%.

Upon release, the residual stress will cause the bimorph beams to bend, holding the

mirror plate out of plane, and parallel to the substrate. The initial elevation is about 600

µm, leaving enough space for the mirror plate to either move downwards or tilt. This

(a)

Mirror Plate

ACT 1

ACT 3

AC

T 2

AC

T 4

Bonding

pads A B

Mirror plate

Bimorph III

(l3) Frame II (L2)

Bimorph II (l2)

θ0

θ0

Frame I (L1)

Bimorph I (l1)

67

mirror can realize piston motion, when the same voltages are applied on all four

actuators; it can realize tip-tilt motion, when differential voltages are applied on 2

actuators on opposite sides of the mirror. 2D scanning can easily be performed when 2

sets of differential voltages are applied on the 2 pairs of actuators controlling the tip-tilt

motion in 2 directions that are perpendicular to each other. During scanning, the

effective optical aperture size will not change, because the rotation center is fixed along

the center of the mirror, optical alignment is also made easy as a result.

Figure 3-5. Electron micrographs of the 2D MEMS mirror. A) the whole device, B) the LSF-LVD actuator, and C) the first bimorph section.

Figure 3-6 A and B show the measured characteristics of the MEMS mirror. As we

can see from Figure 3-6 A, the piston displacement response is linear between 1.5 V

and 4.5 V, and more than 600 µm displacement can be achieved with only 5.5 V

voltage. Figure 3-6 B shows that ±31 optical scan angles can be obtained at a

maximum of 5.5 V, of which, ±(5~31°) optical scanning ranges are linear. As shown in

Figure 3-6 C, the average deviation of the measured results from the linear fit is only

about 0.83% for a large optical scan range up to 26. The frequency response of mirror

is shown in Figure 3-6 D. The mirror has a 10 dB cutoff frequency of 13 Hz. Figure 3-6

1mm 500 µm

100 µm

A B

C

68

E and F are two scanning patterns obtained by driving the MEMS mirror. The raster

scan pattern is obtained by differentially driving two opposite actuators with ramp

waveforms of 0.5~3.8 V at 320 Hz (~30) and 0.5~3.8 V at 8 Hz (~36) for the two

orthogonal directions, respectively. The Lissajous pattern is obtained by shifting the fast

axis frequency from 320 Hz to 300 Hz.

Figure 3-6. MEMS mirror experimental results. A) vertical displacement versus voltage, B) optical scan angles versus voltage applied on each actuator, C) linear fit of measured optical angle by driving actuator 1, D) frequency response of the MEMS mirror, E) a raster scan pattern, and F) a Lissaious scan pattern.

E F

C

A B

D

69

Light Weight Large Aperture LSF-LVD MEMS Mirror

Built upon the LSF-LVD actuator design, a light weight large aperture mirror

design was proposed. Having a large mirror aperture has a few benefits, including

making optical alignment easier, and being able to accommodate larger optical beam,

which could in turn help improve the system resolution. A large mirror plate of 3 mm × 3

mm is chosen for this design.

Figure 3-7. MEMS mirror used for lateral scan. A) SEM of the front side of the MEMS mirror, B) SEM of the back side of the MEMS mirror showing the Si ribs structure, C) layout of the MEMS mirror, and D) characterization of the MEMS mirror: optical scan angle vs. voltage.

To be able to support a larger mirror plate, more actuators were employed. Instead

of using one actuator on each side, four identical LSF-LVD actuators were placed on

A B

C D

70

each side, resulting in a total of 16 actuators. However, large mirror plate leads to low

resonant frequency. In order to keep up the bandwidth of the mirror, a light-weight

mirror structure was employed to reduce the mass of the mirror plate. The light-weight

structure is realized by employing silicon ribs on the backside of the mirror plate for

support, instead of using a solid silicon layer. The mirror plate mass was reduced to

21.8% of its original mass. And the measured tip-tilt resonant frequency was 445 Hz.

The MEMS mirror can scan ±6.5° when a 4 Vdc driving voltage is applied. This design

was later used in the free space OCT setup for brain tissue refractive index

measurements.

FDSB MEMS Mirror

To further increase the mirror fill-factor and enable smaller probe designs, a flip-

chip bonding hidden actuator design is proposed. This design is the 4th generation; it

employs a folded-dual-S-shaped bimorph (FDSB) actuator design to compensate the

curling of one bimorph beam and realize vertical displacement [141], as shown in Figure

3-8. Two bimorph beams of equal but opposite curling angles are connected in series to

compensate the curling of each other, therefore pure vertical displacement can be

realized at the end without lateral shift and tilting. For the two bimorph beams to have

the same curling angle, several geometrical requirements need to be satisfied, including

a certain length ratio of the two beams 31 / LL and an optimized thickness ratio between

each layer in fabrication. Another constraint is that for the hidden actuator design, the

total length of each actuator must fit under the mirror aperture, which sets a total length

limit for the bimorph beams. Table3-1lists the optimized structure parameters that yield

an initial displacement of 315 μm by thermal stress at 300°C.

71

Figure 3-8. FDSB actuator design. A) single S-shaped bimorph, B) use 2 S-shaped to convert curling into vertical displacement, and C) actual FSDB actuator structure.

Figure 3-9. FDSB MEMS mirror. A)schematic with mirror surface facing down, B) schematic with mirror surface facing up, C) SEM with mirror surface facing down, and D) SEM with mirror surface facing up.

B

Bimorph beams

Mirror surface

Flex-PCB

A

C D

SiO2 Al A B C

72

Table 3-1: Optimized structure parameters of FDSB actuator

Fixed Values Optimized Values

Thickness (µm) Adhesion SiO2: 0.05 Function Al : 1.0 Function SiO2 (1st): 1.0 Heater Pt: 0.2 Insulation SiO2: 0.2 Function SiO2 (2nd): 1.4 Adhesion Cr: 0.01

Section length (µm) Total length L: 410 L1=252, L2=20, L3=138

Figure 3-10. Characterization of the FDSB mirror. A) DC response of vertical displacement, B) DC response of optical angle, and C) frequency response.

A

B

C

73

This mirror uses the device layer of SOI wafer as the mirror plate, and the

actuators are formed on top of the wafer. After the device is released, the mirror plate

move downwards and got hidden inside the device. The electrical connecting pads are

formed on the surface of the device frame, and the mirror surface is on the other side.

When the entire device is flipped over, the MEMS mirror can be used as a scanning

engine. The schematic and SEM picture of this design is shown in Figure 3-9.

The mirror aperture size is 1 mm × 1 mm, and the chip size is 1.55 × 1.7 × 0.5

mm3. The fill-factor of this design is as high as 32%, and the device volume is only 1.3

mm3. An initial downward displacement of 340 µm was achieved, and after a self-

annealing and burn-in process, the displacement can increase to 375 µm. The mirror

plate is 25 µm below the top surface when the mirror is flipped over for scanning.

Characterizations of the mirror are shown in Figure 3-10. Applying 4.8 Vdc on all four

actuators, vertical displacement of 319 µm has been achieved, and by driving opposing

actuators differentially, optical scanning angle up to ±23° has been obtained. The

frequency response figure suggests that the mirror has a resonant frequency of about

406 Hz, and the thermal decay is only 1 dB at 200 Hz. The mirror surface curvature is

about 10 m, and the surface roughness is ~20 nm.

Summary

This chapter first introduced the principle of electrothermal actuation mechanism.

And then four generations of MEMS mirrors developed at the BML were presented. The

first generation MEMS mirrors are capable of realizing large scan angles of up to 40°.

However, they have the disadvantage of mirror center shifting during scanning. The

second generation MEMS mirrors overcome the center shifting problem, and can realize

large piston motion movement. Yet, they could not realize tip-tilt motion. The third

74

generation MEMS mirrors employed a LSF-LVD actuator design. And they could realize

both tip-tilt and piston motions, while maintaining mirror center position during scanning.

Detailed designs and characterizations of two types of the third generation MEMS

mirrors were introduced. The first type employed four LSF-LVD actuators, and achieved

25% fill factor, with a mirror aperture size of 1 mm × 1mm on a mirror chip of 2 mm × 2

mm. This mirror design was used for the first three generations of endoscopic probes.

The other type of the third generation mirrors is the light weight large aperture MEMS

mirrors. They employed 16 LSF-LVD actuators, placing 4 on each side of the mirror

plate, and realized a mirror aperture size of 3 mm × 3 mm. These mirrors were used in

the free space OCT system for refractive index measurements. The fourth generation of

MEMS mirrors achieved an even higher fill factor of 32%, by decreasing mirror chip size

to 1.55 mm × 1.7 mm, while maintaining the large mirror aperture size of 1 mm × 1mm.

Design and characterization of these mirrors were also introduced in details. This mirror

design was employed in the fourth generation MEMS probes.

75

CHAPTER 4 OCT ENDOSCOPIC IMAGING BASED ON MEMS

This chapter focuses on endoscopic OCT imaging using probes built with MEMS

mirrors. Four generations of probes will be introduced. The optical and mechanical

designs of the probes will be presented first; then the probe assembly and packaging

schemes will be discussed; after that, results of probe in vivo OCT imaging experiments

will be demonstrated; issues with probe imaging will be discussed at last.

MEMS Endoscopic Probe Design

The goal of this study was to build a side-view endoscopic imaging probe using

the electrothermal MEMS mirrors as the scanning engines. The endoscopic probes

could be used by themselves for OCT imaging or be fitted into an existing endoscopic

probe so that other functions could be performed as well as OCT imaging. The desired

probe size is less than 2.8 mm. To reach this goal, four generations of probes were

built, the packaging schemes were improved along the way, and the probe sizes were

reduced from 5 mm to 2.5 mm by the 4th generation.

OCT probes are connected to the sample arm of the OCT systems, they serve

several purposes. They direct light onto the target sample, and collect the signal back-

scattered from the sample and transfer it back to the system to form interference

patterns. The probe should be able to focus the beam from sample arm to 1-2 mm

underneath the tissue surface, and scan the incident beam transversely to form 2D or

3D images. Figure 4-1demonstrates a typical OCT probe. It is consisted of four

components: optical fiber, graded index (GRIN) lens, MEMS mirror and the probe head

mount that holds the other 3 components. In the following sections, detailed optical and

mechanical designs will be presented.

76

Figure 4-1. Typical MEMS probe design.

Probe Optical Design

The three major optical components in an OCT probe are single mode optical

fiber, GRIN lens and MEMS scanning mirror. Optical fiber is used to deliver and collect

optical signal from and to the system; GRIN lens is used to focus the light beam; and

MEMS mirror is used to direct the focused beam for transverse scans. Figure 4-2

demonstrates the positional relationship of the components (optical fiber is omitted; only

the output beam from the fiber is shown in the figure).

Figure 4-2. Optical components of the probe.

MEMS mirror

GRIN lens

Output light beam from fiber

77

Single mode fiber (corning SMF-28) is chosen to avoid modal dispersion in light

transmission. Core diameter of the fiber is 8.2 µm, its effective refractive index is 1.4677

at 1310 nm, its numerical aperture is 0.14, its cladding diameter is 125±0.7 µm, and the

coating diameter is 900 µm. One end of the fiber has FC/APC connector to connect with

the OCT system, while the other end (distal end) emits light to the GRIN lens, and the

connection is made by using optical adhesive (Norland Optical Adhesive 61), when

cured, the refractive index of the adhesive is 1.542 at 1310 nm. Here, the optical

adhesive not only connects the fiber and GRIN lens physically, it also helps to reduce

refractive index mismatch. The distal end of the optical fiber is cut with an 8° angle to

avoid backreflection from the fiber end, which, if not removed, may saturate the OCT

signal detected by the photodetector. GRIN lens used for focusing is of cylindrical

shape, which, compared to traditional circularly shaped lenses, is easier to package in

smaller spaces. The GRIN lens has a center axis refractive index of 1.6164, a gradient

constant of 0.8521 mm-1, and a working distance of about 5 mm for 1310 nm when it is

placed about 0.034 mm after the fiber. The diameter of the GRIN lens is 700 µm, and

the length is about 1.955 mm. The MEMS mirror is placed at about 1-2 mm after the

GRIN lens. The exact distance varies with different mechanical designs, which is

affected by probe diameters and working distances desired, and more details will be

discussed in the following section.

Probe Mechanical Design

Mechanical design of the mount base is important, because the mount base is

used not only to hold the components in place, but provide electrical connection to the

MEMS mirror.

78

The first 3 generations of endoscopic OCT probes were designed and assembled

based on the 2 × 2 × 0.5 mm3 LSF-LVD mirrors. The first 2 generations are designed to

have a diameter of 5 mm and a working distance of about 1.5 mm in air, as shown in

Figure 4-3. A 0.5 mm-deep cavity is cut out in the base for the MEMS chip to fit in, and

the mirror plate is 45° to the central axis of the probe.

Figure 4-3. Mechanical dimensions of the probe (not drawn to scale).

For the 1st generation probe, a two-piece design using Lucite as material was

chosen. One piece is used to hold the MEMS mirror, while the other piece is for the

other optical components, including fiber and GRIN lens. The MEMS piece has a

protruding part, and the other piece has a matching slot, through which, the two pieces

could be registered together. The two-piece design is intended to protect the MEMS

mirror by separating its assembly process from that of the other optical components,

therefore increasing the chance for successful assemblies. Electrical connection was

made through wire-bonding and conductive gluing. First, five gold wires were bonded to

the pads on the MEMS mirror chip. Then, the MEMS chip was placed and glued in the

cavity designed for it. Next step was to glue the gold bonding wires to the copper wire

embedded in the probe using silver epoxy (Figure 4-4 B). Then two pieces were

79

registered together through the slot. After all the above steps were completed, the probe

was inserted into flexible, biocompatible, transparent fluorinated ethylene propylene

(FEP) tubing with 5.8 mm outer diameter. A finished probe is shown in Figure 4-4 C.

Figure 4-4. 1st generation probe design. A) fabricated probe body in Lucite, B) electrical connection of MEMS mirror to the probe, and C) packaged probe next to a one cent coin.

There are several issues associated with this design. First, the two-piece design

increases the requirements for fabrication precision in order to obtain desired alignment

between the mirror and the optical components. Second, the process of embedding

copper wires increases the fabrication complexity results in price rise of the probe.

Third, high temperature required for curing silver epoxy can cause the Lucite material to

deform, which will cause difficulty, sometimes even failure in registering the two parts

during assembly.

To overcome these problems, a 2nd generation design using one piece base

mount was proposed, alignment was made easy using this new structure. Stainless

steel is chosen to be the probe material since precise fabrication of this material is

readily available, and its deformation under high temperature is negligible. Copper wires

are not embedded in the probe; rather, they are connected through a rigid custom-

designed printed circuit board (PCB). The PCB was designed to fit into the slot on the

A B C

80

top of the probe. On the PCB, five through holes (vias) were drilled to provide

connections from the front layer to the bottom layer; on the front layer, five conductive

pads were placed in a row, each connected with one via through a thin copper wire

embedded in the layer. To form electrical connection, five copper wires were put

through the vias and soldered from the back side of the PCB; and the bonding wires

from the MEMS mirror are glued to the pads on the front layer of the PCB. The rest of

the assembly process is similar to the design before. The diameter of the probe mount

was 5 mm; after being inserted into the FEP tubing, the packaged diameter was 5.8mm.

Figure 4-5. 2nd generation probe design. A) 3D model of the assembled probe, B) assembled probe in stainless steel.

The successful implementations of the previous designs lead to further probe

miniaturization. To save space in electrical connection, flexible PCB was introduced. 5

copper pads were placed in a row on the front side of the PCB, and these pads are

connected to the back side to 5 soldering pads, which will then be soldered to five

copper wires for driving signals. In this design, a rectangular slot is cut through the

probe, and flexible PCB was inserted through it and glued to the mount base. MEMS

mirror is then placed on top of the PCB. 5 gold wires with one side wire-bonded to the

MEMS mirror and the other side conductively glued to the PCB were used to electrically

A B

81

connection the MEMS mirror to the PCB. This design belongs to the 3rd generation, it

has successfully reduced the diameter of the probe to 2.8mm. 3D views of the probe

are demonstrated in Figure 4-6.

Figure 4-6. 3rd generation probe design. A) mechanical design for the probe, B) electrical connections of MEMS mirror and flex PCB, C) 3D model of the probe, and D) assembled probe with flexible PCB.

Inspired by the 3rd generation design, a new 4th generation prototype based on the

DFSB mirror was developed to further decrease the probe size. This design has a

diameter of 2.5 mm and length of 10 mm. The working distance is designed to be about

2 mm in air. The electrical connection has been made easy due to the special design of

DFSB mirrors, which have their connection pads on the backside of the mirror frame. In

A

B

C D

82

this design, no wire-bonding process is necessary, therefore, a lot of space could be

saved and further miniaturization was possible.

Figure 4-7. DFSB probe design. A) mechanical design of the probe, B) 3D model of probe frame, and C) 3D model of assembled probe.

Similar to the 3rd generation design, this probe uses flexible PCB for electrical

connection. The difference is that electrical connection can be made by directly placing

the MEMS mirror on top of the flexible PCB, therefore eliminating the wire-bonding

process. The pads pattern of the PCB is designed to match those of the mirror, as

shown in Figure 4-8 B. The flexible PCB is first glued to the probe, and then small

amount of silver epoxy is dispensed on the PCB bonding pads, MEMS mirror is aligned

to the PCB and both electrically and physically glued to the PCB through silver epoxy.

After the MEMS mirror is assembled to the probe, GRIN lens and fiber are installed. The

B C

MEMS mirror

GRIN lens slot

Open sidewall for MEMS mirror

Optical fiber slot Flex-PCB

Optical fiber

GRIN lens

Output beam

A

83

probe is then packaged with a transparent FEP tubing, which has an inner diameter of

2.5 mm and an outer diameter of 2.8 mm.

Figure 4-8. DFSB probe design. A) flexible PCB design, B) zoomed in view of the PCB head with bonding pads pattern, and C) packaged probe.

In Vivo Experiments

The 2nd and 4th generations have been employed in in vivo experiments to

demonstrate their abilities. For the two imaging sessions, the samples were the same

type female athymic (nu/nu) nude mouse with a body weight of 20 g to 26 g (Harlan

Laboratories, Indianapolis, IN). The mice were anesthetized by injecting ketamine (100

mg/kg) and xylazine (10 mg/kg) intraperitoneally. 2D transverse scans are done by

simultaneously driving all four actuators with one opposing pair for fast scan and the

other pair for slow scan.

With the 2nd generation probe, the actuator pair on the circumferential direction

was used for fast-axis scan, driven by 0~4 V differential ramp voltages at 1.25 Hz; and

the actuator pair on the longitudinal direction was used for slow-axis scan, driven by

0.5~3.5 V ramp voltages at 0.0125 Hz, therefore accumulating 100 frames of 2D images

Bonding pads Soldering pads

A

B C

MEMS mirror

GRIN lens

Optical fiber

Flex PCB

84

in 40 s for one 3D reconstruction. This generates a transverse area of ~2.3 × 2.3 mm2.

Both 2D and 3D images of the mouse tongue have been obtained. The 2D OCT image

shown in Figure 4-9 A was acquired using the MEMS OCT probe at the base of the

tongue. A stratified squamous keratinized epithelium (SSKE) and basement membrane

(BM) corresponding to the relatively strong signal bands from the upper layers of the

tissue structure were observed. Figure 4-9 B shows a reconstructed 3D image of the

same mouse tongue. 2D and 3D OCT images of the mouse ear have also been

obtained, and the results are shown in Figure 4-10. The ear thickness is about 500 µm.

The dermal structures and the subcutaneously adjacent layers were observed using the

MEMS OCT probe in this experiment. The cartilage (C) of the ear was represented in

the middle by the dark band and the dense conjunctive capsule (cc) with two bright

layers put around the cartilage. The superficial epidermis (E) of the mouse ear was

detected at the first and last of the mouse ear longitudinally. The lower and upper areas

from the dense conjunctive capsule to epidermis, the dermis (D), were observed. Figure

4-10 B shows the 3D reconstruction image of the ear.

Figure 4-9. In vivo OCT images of mouse tongue. A) 2D image of a mouse tongue, B) 3D image of a mouse tongue.

1mm A B

SSKE

BM

SSKE

BM

85

Figure 4-10. In vivo images of mouse ear. A) 2D image of a mouse ear, B) 3D image of a mouse ear.

Using the 4th generation probe, mouse ear and tumor were imaged. For this

imaging session, human breast cancer cells, 4T1 mammary carcinoma cell line, were

injected into the subcutaneous tissue of the female athymic (nu/nu) nude mouse with a

total volume of 10 μl at a concentration of 10,000 cells / 10 μl. And the tumor cells had

been cultivated for ten days by the time of the imaging experiment. The ear imaging

results are shown in Figure 4-11 C. As have been reported for the 2nd generation probe,

the cartilage (C), the dense conjunctive capsule (cc), the superficial epidermis (E) and

the dermis (D) of mouse ear were also observed using this probe. The tumor imaging

results are shown in Figure 4-11 F-J. F and G are 3D reconstructions of the tumor area

from different angles, and H-I are the 2D cross-sectional images taken in areas pointed

in G. The tumor edge can be clearly seen in 3D images, and 2D images demonstrated

obvious difference between normal and tumor tissue areas, due to strong penetration

loss inside the tumor cell.

E

D

E

D

cc

cc

C

1mm A B

86

Figure 4-11. In vivo imaging results. A) mouse under ear imaging session, B) 2D image of mouse ear, C) 3D image of mouse ear, D) probe position of tumor imaging, E) injected tumor under microscope, F-G) 3D mouse tumor imaging results, and H-J) 2D images of tumor margin, locations are labeled in G.

Issues with Probe Imaging

There are a few factors that affect the optical qualities of the imaging probes,

including the relative positions change between the optical components due to the

variation in the assembly process, the mirror curvature, the plastic tubing protecting the

probe and effect of the non-telecentric scan, circled in Figure 4-12. To understand these

issues, Code V (Optical Research Associates) simulations were conducted. And a “free

space probe” was setup to confirm the simulation results.

(e)

1 mm Probe tubing sidewall

E

C cc

cc D C

E

D

A

Probe

Tumor region

Imaged margin

Tumor region

1 mm 1 mm 1 mm

B

C

F

D E

G

H I J

H I J

87

Figure 4-12. Demonstration of issues focused on in this study.

Code V builds optical models by defining properties of each surface. Figure 4-13 A

defines the optical system of the endoscopic MEMS OCT probe. The object surface in

this system is the fiber output; the object beam angle was defined by system numerical

aperture, which is 0.14 for single mode fiber. The distance between the fiber output and

the GRIN lens was defined in the thickness column, 0.05 mm in Figure 4-13 A. The next

two surfaces define the GRIN lens; thickness of 1.955 mm is the length of the lens and

“GRIN1310” is a user-defined GRIN material with center axis refractive index of 1.6164,

a gradient constant of 0.8521 mm-1 as the ones used in the probes. Surface #3 and #4

represent the MEMS mirror. The MEMS mirror is the only reflective surface in the

system, and it has a 45° decenter from the center axis. For the accuracy of simulation, a

dummy surface (surface #3) was used to generate the decenter. The mirror is placed

1.5 mm behind the GRIN lens. FEP tubing was simulated with two cylindrical shaped

surfaces with a uniform refractive index of 1.5; the inner surface diameter is 2.5 mm and

the outer surface diameter is 2.8 mm. The thickness of the image surface was defined

as a variable, meaning it could be changed automatically during simulation to find the

best focus position. In the simulation, the best focus position was found by geometrical

ray tracing. After the system is fully defined, it could be visualized with a 3D drawing as

shown in Figure 4-13 B. In this system, the Gaussian spot size at each plane could be

88

calculated using Gaussian beam tracking. Surface properties and thicknesses could be

varied to study the different issues mentioned earlier.

Figure 4-13. Code V simulation. A) optical system definition in Code V, B) optical system 3D model demonstration.

To confirm the simulation results, a setup was built using 3D movement stages to

simulate a free space probe. The setup consists of three 3D stages, each one

controlling the position of the fiber, the GRIN lens and the imaging target. The target

used in this study was a glass slide with a 1951 USAF resolution pattern. The resolution

of the system was determined by the finest pattern distinguishable from the OCT

images. And the focal length was recorded as the center of the clearest image range.

GRIN lens

Fiber output

FEP tubing

MEMS mirror 50 µm

B

A

89

Figure 4-14. Optical quality testing experiments. A) experimental setup for testing, B) 2D image of distinguishable pattern, and C) 3D image of distinguishable resolution pattern.

Effect of the Distance between the Fiber and the GRIN Lens

First of all, effect of the distance between the fiber and the GRIN lens on the focus

point of the GRIN lens was studied. The distance between the fiber and the GRIN lens

(L1) determines how far away the output light spot from the fiber tip will be focused at

(L2), therefore affects the working distance of the probe. L1 also has an effect on the

size of the focused spot. To understand how much effect the variation of L1 has on L2

and the spot size, code V simulations and experiments were conducted, and their

results showed good agreement with each other.

The results shown in Figure 4-15 suggested L2 changes non-linearly with L1; and

when L1 changes within 0.01 to 0.10 mm, the focal length changes more than 3 mm.

The desired working distance of about 5 mm can be obtained when L1 is controlled at

approximately 0.034 mm; and as long as L1 is smaller than 0.08 mm, the probe will

have an acceptable focal length of larger than 3.5 mm, meaning the beam focus will be

outside the probe sidewall, and could be focused inside the imaging sample. There is a

tradeoff between the working distance and spot size, as the working distance increase,

the spot size increases too. We have found the experimental results are in good

A

GRIN lens

MEMS mirror

Fiber

Target

B C

49.6 µm

90

agreement with simulation results when spot size is at 60% of the maximum intensity

from the simulation. The spot size is larger than 30 µm within the acceptable working

distances for this probe design.

Figure 4-15. Effects of the distance between the fiber and the GRIN lens. A) schematic of the testing setup, B) effect on focal length, and C) effect on spot size.

C

B

A

91

Effect of Mirror Curvature

Due to the fabrication process, MEMS mirror surface is not perfectly flat, and the

radius of curvature of the mirror not only affects the focal point position of the probe, but

generates astigmatism in the system, which could cause imaging quality to degrade. In

Code V simulation, Y radius for surface #3 controls the mirror curvature (Figure 4-16 A);

by changing this value to different radii of curvatures, different optical quality

corresponding to different mirror curvatures could be evaluated. Several mirror radii of

curvatures reported by our group was chosen for simulation. The smallest radius of

curvature reported was 0.132 m, and improved radii of curvatures of 0.15 m, 0.18 m, 0.8

m and up to 10 m for the 4th generation mirror were reported. In addition to these

reported values, two extreme cases of 0.1m and infinity (perfectly flat mirror plate)

radius of curvature were also simulated.

Figure 4-16. Code V simulation for effect of mirror curvature. A) optical system definition, B) 2D view of the optical system shown with 5 mm mirror curvature.

A

B

A

B

92

Figure 4-17 was obtained at L1 = 0.05mm. Focal length decreased with the

increase of mirror radius of curvature. This can be easily explained by the converging

capability of concave mirrors. Light converges faster with mirrors that have smaller

radius of curvature. And because of the same reasons, the focused spot size is smaller

when radius of curvature is smaller; this may seem desirable for the sake of higher

resolution; however, the drawback is that astigmatism becomes more and more serious

as the radius of curvature decreases. When there is astigmatism in the system, light in

the sagittal direction and the meridional direction will be focused at two different planes,

the focus for this type of system is then in between the two focal planes where the circle

of least confusion is formed. As a result of astigmatism, the focused spot deform to an

elliptical shape, which would adversely affect the OCT image quality. From Figure 4-17

B, the eccentricity of spots remains smaller than 0.14 when the radius of curvature is

larger than 500 mm. And at 10 m radius of curvature, the astigmatism in the system

could be neglected.

These simulation results were compared with the focal lengths and resolutions

tested with the free space setup. The MEMS mirror used for this experiment is the 4th

generation FDSB MEMS mirror, and the mirror radius of curvature is about 10 m. Figure

4-18 A is the comparison between the tested focal length and the simulation results

when mirror curvature is 10 m. And Figure 4-18 B compares the tested system

resolution when L1 is 0.05 mm and 0.1 mm to the simulated full width half maximum

(FWHM) spot size. The results have suggested the tested result agrees well with the

simulation results for focal length. For the resolution, we can see that we could

distinguish slightly better than FWHM of the spot diameter with the MEMS scan system.

93

Figure 4-17. Effects of the MEMS mirror radius of curvature based on simulation. A) radius of curvature vs. focal length, B) radius of curvature vs. spot size.

Figure 4-18. Comparison of simulation and experimental results. A) distance between fiber and GRIN lens vs. focal length, B) distance between fiber and GRIN lens vs. spot size (simulation result is FWHM spot size at 10m radius of curvature, experimental result is best observed resolution pattern size).

Effect of FEP Tubing

In order to protect the probe head, after all the assembly process is done, the

probe is inserted into a fluorinated ethylene propylene (FEP) tubing. This section

focuses on the effects of this tubing on the optical qualities of the system.

First, to understand how adding the tubing would affect a system with no existing

astigmatism, MEMS mirror was excluded from the simulation; the tubing was placed

A B

A B

94

directly after the GRIN lens, as shown in Figure 4-19 A. This FEP tubing has an effect

similar to that of a cylindrical lens, introduces astigmatism to the system. From the result

shown in Figure 4-19, with the FEP tubing, focal length change increased from 0.0357

mm (at L1 = 0.2 mm) to 0.3561 mm (at L1 = 0 mm). At L1 = 0.34 mm, when the focal

length is about 5.0219 mm without FEP tubing, focal length increases to 5.0721 mm

when tubing is added. The increase is a result of the higher refractive index of the FEP

material, as demonstrated in Figure 4-19 D. As a result of astigmatism, at the focal

plane, the spots deform to elliptical shape. Figure 4-19 C shows the X, Y spots

dimensions at the focal point with different L1. For the probe to have a usable working

distance, the focal length should be larger than 3.5 mm; under these circumstances, the

eccentricity of spots are smaller than 0.24, and spot radius increased by about 1.5 µm

at most.

Then, the MEMS mirror was added back to the system simulation. As shown in the

last section, when mirror is not perfectly flat, it also causes astigmatism in the system. In

this case, the effect of mirror curvature and FEP tubing combines, and simulation result

have shown increased spot size and larger eccentricity at the same mirror curvature.

Figure 4-20 A shows that spot size increases more in the Y direction, and the spot size

increase is slight more evident for larger mirror curvature. Focal lengths increased

almost uniformly with FEP tubing. Using flat tubing or astigmatism correction lens in

combination with the cylindrical tubing may reduce the astigmatism in the system, and

may be further studies in the future.

95

Figure 4-19. Effects of FEP tubing on optical beam from the GRIN lens. A) schematic of the simulation, B) effect on focal length, C) effect on spot size, and D) demonstration of light travel through higher index material.

Y

X

A

B

C

D

96

Figure 4-20. Effects of FEP tubing on the system. A) effect on spot size, B) effect on focal length.

Effect of Non-Telecentric Scan

The last issue we consider is the non-telecentric scan effect of the probe. Since

the scanning mirror is placed after the GRIN lens, imaging plane is not flat. There are

two scan directions, longitudinal and radial; Figure 4-21 demonstrates the non-

telecentric scan effect in the longitudinal direction.

From basic geometrical optics, when the incident beam and the normal to the

reflection plane are in the same plane, reflection beam will be in the same plane; and

A

B

97

when the mirror rotates in plane, the angle change of the reflection beam will be twice of

the rotational angle of the reflection plane. In the longitudinal scan direction, when the

MEMS mirror scans linearly, the optical beam will also scan linearly at twice the angular

speed. According to the simulation, as long as the mirror plate is flat, the change in spot

size is very small at different scan angles. However, it does cause some deformation in

the OCT image. For each cross-sectional scan, a row of pixels shows information at the

same optical depth; as a result, a flat surface would be shown with a curvature in the

image, as shown in Figure 4-22. A simple algorithm could be carried out to improve the

visualization, which will be described in the following. For each A-line signal collected

during scanning, its offset from the center position was calculated and then entire A-line

was shifted accordingly to match up with the center position. Specifically, in the OCT

system, controlling signal for the MEMS mirror is outputted through a data acquisition

card (USB-1408FS, Measurement Computing). The output signal of 0-4 V was divided

into 112 steps and output over 0.4 s, 512 A-lines and 4096 points for each A-line were

collected during this time. To obtain the image in Figure 4-22, scanning range of ±12.5°

was measured at about 2 mm from the MEMS center, the entire imaging depth was 1.6

mm. Therefore, for the Kth A-line, the number of points need to be moved could be

calculated by equation (4-1). The initial result is shown in Figure 4-22 B. This algorithm

greatly simplified the problem by moving each A-line all together by the same amount, a

more accurate algorithm that takes into account the scanning angle is desirable for

more accurate display.

mmkmm 6.14096))512

1112

1112

255.12cos(1(2

(4-1)

98

Figure 4-21. Demonstration of the non-telecentric scan effect. A) imaging area during scan, B) imaging plane direction at a certain scan angle.

Figure 4-22. Correction for curved image display. A) 2D image of a flat IR card before correction, B) 2D image of the flat IR card image after correction.

The scan in the radial direction is a little more complicated. As shown in Figure 4-

23, for every different angle in radial direction, the input beam and the normal to the

mirror plate would form a different plane, and the law of reflection will take place in the

new plane. As a result, when the mirror scans in the radial direction, the reflected beam

would form a curved imaging plane. Using space geometry, the relationship between

the mirror scan angle β in the radial direction and the beam scan angle θ could be

calculated by equation )2

cosarccos(

. A non-linear relationship is suggested by this

equation. 3D reconstruction algorithm is needed in order to correct the distortion in this

direction and this will be a research focus in future studies.

A B

Imaging area

offset

A B

99

Figure 4-23. Non-telecentric scan effect in radial direction. A) with one scan angle, B) with multiple scan angles.

Summary

In this chapter, four generations of MEMS based endoscopic OCT probes were

introduced. The probe diameter reduced from 5 mm to 2.5 mm by the fourth generation.

During this process, the packaging schemes were simplified with the introduction of

flexible PCB and a novel mirror design that places the connecting pads on the bottom of

the mirror chip. Capabilities of the probes were demonstrated with in vivo imaging

experiments. Clear layered structures and tumor shape were visible from the OCT

images obtained with the packaged probes. At last, four factors that affect the

performance of the probes were studied. We could conclude that by precisely

controlling the distance between the fiber and the GRIN lens, an acceptable focusing

length range of larger than 3.5 mm could be obtained, with lateral spot size slightly

larger than 30 µm. The working distance and resolution were not significantly affected

by the 4th generation MEMS mirror, due to its large radius of curvature. FEP tubing

introduces slight astigmatism, when MEMS mirror is not perfectly flat. And the resolution

will not change much at different scan angles during imaging.

A B

100

CHAPTER 5 MEMS BASED FREE-SPACE TISSUE IMAGING — REFRACTIVE INDEX

MEASUREMENT

In this chapter, we explore another MEMS based OCT application in free space:

localized refractive index (RI) measurement. In this study, OCT is used to measure

group RIs in various anatomical regions in live rat brain tissue slices. The effects of

tissue deformation on RI changes are also measured in different anatomical regions. To

measure RI, OCT and an indenter tip are used to measure both the physical and optical

thicknesses of normal and compressed tissue slices, and cell viability for tissue slices is

maintained throughout the measurements. RI is determined by the ratio of physical

thickness and optical thickness of the brain tissue slices. The large aperture LSF-LVD

MEMS mirror was employed for transverse scan in order to minimize disturbance to the

test sample.

RI Measurement Background and Motivation

The refractive index (RI) of a sample measures how fast light travels in it. Even

with high resolution imaging modality, optical images of a biological tissue can still be

distorted by inaccurate data of the tissue’s RI, which may vary from region to region

even in the same tissue. Variation in RI between anatomical regions may be due to

heterogeneous tissue structures caused by differences in tissue constituents. Variation

in RI may also be introduced with compression or swelling of tissues, which can change

the ratio of solid and fluid fractions in tissue. Accurately measuring RI is essential to

provide high-fidelity optical images since RI is a key parameter for image reconstruction

and interpretation as well as for understanding light-tissue interactions. Besides, clinical

diagnosis can be made based on the refractive index difference between normal and

malignant tissues [142], [143].

101

In previous studies, numerical simulations based on photon absorption and

scattering models have been used to predict tissue RIs [143], [144]. These methods

involve complex algorithms and lengthy calculation times. Other attempts to

experimentally determine tissue RIs measure the output angle of an optical fiber with

the tissue as the fiber cladding [145], or measure the critical angle of total internal

reflection between the boundary of the tissue sample and a triangular prism [146] or

semi-cylindrical lens [147]. Using these experimental methods, the RIs of various

tissues including kidney, liver and striated muscles were obtained [145], [147].

Recently, more accurate methods for measuring RIs were proposed by Tearney et

al. based on OCT [148]. Two methods were introduced to measure RIs of dermis,

epidermis and stratum corneum in [148]: the first one used OCT images of a flat

substrate and tissue boundaries and extracted the RI by comparing the optical and

physical thickness; and the second method utilized focus tracking, in which the

movements of the reference and sample arms were recorded and the RI was calculated

based on the distances they moved. Binding et al. measured the RI of the cortex of the

rat brain as a function of age by employing a full-field OCT with a similar focus tracking

method [149]. Note that a broadband light source is typically used in OCT, so group RI

was measured in these methods. Another approach, which was applied to liver, spleen

and brain tissues, measured the phase change induced by tissue samples to calculate

RI [150]-[152]. This method, by using a HeNe laser as the light source, gives the phase

RIs of the samples. Group and phase RI are defined with respect to group and phase

velocity respectively, and they are slightly different in biological tissues due to the

chromatic dispersion.

102

While optical properties have been previously measured in central nervous system

(CNS) tissue, RIs for living tissues are scarce, and to our knowledge, regionally-varying

RIs in brain tissue have not been reported. Measurement of regionally varying RI is

important for heterogeneous tissues consisting of different types of cells and cellular

densities at each anatomic region. With these additional measures, distortion of optical

images between anatomical regions due to variation of RIs may be corrected in

complex tissues. Maintaining tissue viability is potentially important for measuring RI of

CNS tissues. Hypoxic neuronal injury due to a lack of transported oxygen and nutrients

increases cell membrane permeability, and cell membranes and matrix components will

also lose their structural integrity. These physical changes may have an effect on optical

properties of brain tissues. In addition, information on how RI changes with regards to

physical deformation may be used to characterize or potentially diagnose diseases

which are related to tissue deformation such as hydrocephalus and brain tumors [153]-

[155]. Besides, OCT-based elastography, which has been used for medical diagnosis

and mechanical testing may also benefit from this knowledge, since it is based on the

measurements of tissue deformation under loaded conditions [156], [157].

Free Space OCT System

The same time domain OCT system was employed for imaging the optical

thickness of brain tissue slices. The sample arm is set up in free space for this

application. In the sample arm, instead of using a translational stage to hold the sample

for lateral scans, the light weight large aperture LSF-LVD MEMS mirror introduced in

Chapter 3 was used (Figure 3-7). The output light from the single mode fiber (SMF-28)

was collimated with a fiber collimator (Thorlabs, F260FC-C); and the collimated beam

size was 2.8 mm in diameter. The light was then focused using a lens with diameter of

103

12.7 mm and focal length of 39 mm. A MEMS mirror was placed approximately 26 mm

behind the focusing lens, with 45° to the incoming beam. The beam diameter at the

mirror plate was about 0.94 mm, and the 45° cross section was about 1.33 mm in

diameter. The mirror plate is 3 mm × 3 mm, which is larger than twice the beam size

and did not truncate the incident beam even at the largest scan angle of the MEMS

mirror. The transverse resolution was ~23 µm.

Figure 5-1. OCT setup for refractive index measurement.

The advantage of using a MEMS mirror was that it allowed the sample to remain

stationary, and thus disturbance from a moving stage was avoided. For ex vivo testing,

the tissue slices were submerged in a physiological fluid, but were not fixed to the

underlying substrate. With stage movement, even small relative movement of the tissue

would cause image artifacts, and by employing a MEMS mirror, motion artifacts were

avoided.

Sample Preparation

Tissue samples from five adult male Sprague Dawley rats (~ 250 g) were tested.

Tissue slices from three rats were used to measure RIs within different anatomical

regions and tissue slices from two rats were used to measure RI changes with tissue

104

compression. Rat surgery was conducted in accordance with the NIH guidelines on the

use of animals in research and the regulations of the Animal Care and Use Committee

of the University of Florida. Prior to euthanasia, rats were fully anesthetized by

isoflurane inhalation and checked for absence of toe-pinch, righting, and corneal

reflexes. After euthanasia, standard protocols for tissue retrieval, brain tissue slicing,

and tissue maintenance were implemented. Excised rat brains were immediately sliced

using a vibratome (Leica VT 1000A, Leica Microsystems Inc., Germany) into coronal

sections of 300 µm thickness. This thickness was chosen to ensure transport of oxygen

and good visibility through the sample thickness in OCT images [159]. After slicing,

brain tissues were submerged in O2-saturated artificial cerebrospinal fluid (aCSF,

Neurobasal™ Media, GIBCO, Invitrogen Co., CA) and temperature was maintained

between approximately 35 - 37 °C. The aCSF was continuously bubbled with 95% O2

and 5% CO2 gases and 0.5 mM L-Glutamine (Invitrogen Co., CA) and 1% penicillin-

streptomycin (Invitrogen Co., CA) were supplemented. The pH of the aCSF measured

before testing was 7.4. Through testing, sliced brain tissues were maintained under

these conditions; cell viability was determined by measuring neuronal death or

degeneration as a function of incubation time with Fluor-Jade C (FJC) staining. The

results have shown that cells in the slices could maintain viability for up to 7 hrs [157].

Method and Process

To establish regional variation of RI in acute rat brain tissue slices, cerebral cortex,

putamen, hippocampus, thalamus and corpus callosum were tested (Figure 5-2). The RI

of a tissue slice is calculated by dividing optical thickness by its physical thickness, i.e.

physical

optical

t

tRI

(5-1)

105

Thicknesses were measured using the following procedure. To measure RI, an

acute rat tissue slice was placed into a Petri dish which was filled with oxygenated

aCSF (aCSF was oxygenated with 95% O2 and 5% CO2 for an hour prior to testing). A 2

mm spherical bead attached to a micro-positioner was lowered onto the slice to touch

the surface of selected anatomical regions, see Figure 5-3 A. The spherical ball was

aligned within the center of the lateral scan range of the MEMS mirror and OCT was

used to scan cross-sectional tissue images from the bottom of the Petri dish which was

made of glass. The optical thickness of the tissue slice was obtained by directly

measuring the thickness of the slice in the OCT image through the number of pixels

(see Figure 5-3 B). After the optical thickness was measured, the brain tissue slice was

removed gently without moving the metal bead, and the aCSF in the Petri dish was

removed using a pipette and paper tissue. The physical thickness of the brain tissue

slice was determined using the thickness of the air gap between the metal bead and the

bottom of the Petri dish. The RI of air is 1. The RI was calculated by comparing the pixel

numbers of the tissue slice thickness and the air gap. To verify accuracy, a known

optical path length of 500 µm was measured to be between 496 to 498 µm using this

setup, so the system error was smaller than 0.8%. Ten tissue slices from each region

were measured; one test was conducted on each slice.

Figure 5-2. Anatomical regions of rat brain tissue slices tested. Medial sections from excised rat brains were sliced into coronal sections of 300 µm thickness.

106

To measure the refractive index of tissues under compressed states, the spherical

bead was replaced by a 2 × 2 × 0.2 mm flat glass tip to apply uniformly compressive

strains. Brain slices from 2 rats were used for this study, 7 slices for cerebral cortex and

8 slices for corpus callosum were tested. The same area in each slice underwent strains

from 20% to 80% during a single series of tests. The initial physical distance between

the bottom and top surface of brain tissue was measured using a micrometer stage, and

then tissue was deformed by moving the flat indenter. Brain tissue slices were

continuously compressed by 20%, 40%, 60% and 80% with respect to the initial

physical thickness. Cross-sectional deformed images were taken by OCT and the

physical thickness and optical thickness were compared to measure changes of RI with

respect to tissue compression. The experimental procedure took less than 2 minutes for

each series of tests. OCT images for these experiments are shown in Figure 5-5.

Figure 5-3. Measurement process. A) Demonstration of positioning the metal bead during the experiment, B) OCT image of the brain tissue slice, and C) OCT image of air after brain tissue is removed.

A

B

C

tissue thickness

air thickness

107

Figure 5-4. Measurement of optical and physical thickness of various anatomical regions in rat brain tissue slices. A1-E1) meansurement of the optical thicknesses (toptical) of each target region, and A2-E2) measurement of the physical thicknesses (tphysical).

A1 A2

B1 B2

C1 C2

D1 D2

E1 E2

108

Figure 5-5. OCT images of brain tissue slices under compression for A) cerebral cortex (grey matter) and B) corpus callosum (white matter, bundle of myelinated axons). A1, B1) no compression, A2, B2) 20%, A3, B3) 40%, A4, B4) 60%, and A5, B5) 80% compression.

Results

OCT provided 2D cross-sectional images of acute rat brain tissue slices, shown in

Figure 5-4 and Figure 5-5. The bottom and top surfaces of the tissue slices were clearly

delineated. The surface contact of the 2 mm-diameter stainless steel bead was also

detected for each anatomical region. The RI in the corpus callosum, which is a known

white matter region, was 1.407 ± 0.015 and the RIs in the putamen and cortex, which

are known as grey matter regions, were averaged between 1.361 and 1.369. RI in the

corpus callosum was measured to be statistically higher than those in other regions of

brain tissue (see Figure 5-6). The corpus callosum was the only region found to be

statistically different from other anatomical regions based on the Tukey-Kramer method

with alpha=0.05. The standard deviation of the RI in each region ranged from 0.007 to

0.016. The RI of the aCSF was 1.342 ± 0.007. This was ~2.7% lower than the average

RI over all the regions measured in the brain tissue slices, which was 1.380.

A1 200µm

A2 A3 A4

B1

A5

B2 B3 B4 B5

109

Figure 5-6. Measured RI in various anatomical regions in brain and aCSF. 10 samples were measured at each region. Bars correspond to 1SE (CC=corpus callosum, Thal=thalamus, Hip=hippocampus, CCt=cerebral cortex, Put=putamen, and aCSF=artificial cerebral spinal fluid). RI of corpus callosum was statistically different compared to other regions.

The effect of tissue deformation on RI was estimated in the cerebral cortex and

corpus callosum by unconfined uniform compression. The RI changes in both regions

showed a similar trend, non-linear increases with increasing compressive strain, as

shown in Figure 5-7. The RI of the white matter increased more significantly under the

same strain compared to the RI of the grey matter. For strains up to 20%, only small

increases in RI were observed, which were ~3% for the grey matter and 6% for the

white matter. For applied strains over 40%, RI significantly increased. For 80% strain, a

RI increase of more than 70% was measured for the grey matter, while an average 90%

increase was measured for the white matter.

110

Figure 5-7. RI in rat brain tissue slices under uniform compression. 7 rat brain tissue slices were test for cerebral cortex and 8 tissue slices were tested for corpus callosum measurements. Bars correspond to ±1SE.

Discussion and Conclusion

In this study, RI of rat brain tissue slices was measured in various anatomical

regions using a MEMS based OCT system. OCT images clearly captured cross

sectional images of brain tissue slices with white matter (corpus callosum and thalamus)

showing brighter intensity than gray matter regions such as the cerebral cortex and

putamen. RI of the corpus callosum, which consists primarily of white matter was

measured to be 1.407±0.015. No statistical differences in RIs of grey matter regions

(cerebral cortex and putamen) were found, and the average RI of all grey matter regions

was 1.369±0.014. RI of aCSF was slightly higher than RI of water by approximately

0.9%. The difference between RIs at each anatomical region may be due to different

types of cells, cell densities, and differences in fractions of extra/intracellular space in

their structures. The higher RI measured from the white matter may also be due to

multiple scattering. RIs of both white and grey matter regions were close to RI of water,

111

because of the large water content of brain tissue [160]. RIs of brain tissue slices

measured by OCT in this study are comparable to those previously reported in

literature, for example, 1.3526 for cortex at a 1.1 µm wavelength [149], 1.371 over a

range of spatial frequency of 0.06 to 0.6 µm-1 (region not specified) [150]; and 1.3751

and 1.3847 for two cortical neuron cell bodies at 658nm [151]. Differences between

measures may be due to different measuring wavelengths. Another reason could be

due to maintenance of tissue viability. Tissue viability may be important in measures of

optical properties in brain tissue since it can change with the fraction of solid to fluid

constituents, e.g. extracellular fluid content. However, our measured RIs may be similar

to other studies since ~ 80% of brain tissue is fluid, and changes in solid to fluid volume

fractions with changes in tissue viability may not be significant as long as structural

integrity of the ECM and cell membranes is maintained (short time after slicing).

Changes of RIs under unconfined compression testing of brain tissue slices were

also measured in both white (corpus callosum) and grey matter (cerebral cortex)

regions. Non-linear increasing trends with increases in strain were observed for RIs of

both regions. Under smaller compression strains under 20%, changes of RI were almost

negligible. RIs increased significantly over 40% compression. This large increase may

be due to changes in water content and significant compaction of solid matrix in the

tissue with large strains as fluid is squeezed out of the extracellular space. As

compression approached 80%, most extracellular fluid was likely squeezed out and RI

which was significantly higher than for non-compressed states likely reflected that of the

solid constituents including cells, vessel walls, and ECM. The higher scattering of solid

112

constituents compared to fluids can therefore be used to explain the gradual increase in

OCT image intensity with the increased compression shown in Figure 5-5.

In measurement of RIs, tissue swelling can be a significant issue when comparing

acute slices to tissues in vivo. Thicknesses of brain tissue slices were measured within

2 hours after tissue slicing. Average thicknesses of the cerebral cortex and corpus

callosum were 341 µm (1SE=14 µm) and 254 µm (1SE=22 µm), respectively. Thus an

approximately 13% thickness increase in the cerebral cortex and a 20% thickness

decrease in corpus callosum were detected after tissue slicing. This may in part be due

to effects of residual stresses within brain tissues. For example, within the brain, certain

grey matter regions have been measured to be in compression while white matter was

measured to be in tension [161]. After slicing, these stresses may be removed resulting

in thickness changes. However, these changes in the thickness may not have a

significant effect on RI, since changes of RI under 20% of compression were estimated

to be less than 3%. This suggests that RI measurements in vivo and ex vivo could be

comparable.

It should be noted that blood flow was not taken into account for this study, and

the temperature of brain tissue samples may have changed slightly during the testing

process since the sample arm was not heated. These two factors may introduce some

differences when comparing the testing results to in vivo conditions. In large strain tests,

viability of cells may be compromised as tissue is compressed. Within the region of the

flat glass tip, the extracellular space may be greatly reduced, resulting in hindered

extracellular transport and hypoxic cell damage. To reduce this effect, all test

procedures were completed in less than 2 min within each test region. Cell viability may

113

not be significantly compromised during this short period of time. Cells may also be

damaged due to loading under large compression (over 40 to 80% of compression).

With rupture or damage to the cell membrane, cellular components are released and

these constituents may cause further tissue damage. Also due to the small scan angle

of the MEMS mirror, the non-telecentric scan effects were not corrected within the OCT

images. The precision of RI measurements can be increased with improved resolution

of the OCT system and correction of the non-telecentric scan effects.

The free space OCT system combined with indentation technique is simple to

implement, it takes advantage of the high resolution of OCT and the region selection

capability of indentation. The results of this study may be helpful in correcting optical

image distortion caused by regionally varying RI and tissue compression caused by

injury, surgery or disease. Accurate optical images of brain tissue could aid in image

interpretation and may lead to improved diagnosis and more successful imaging-guided

surgery. In addition, OCT-based mechanical testing of tissues under compression or

indentation tests may also benefit from this study.

114

CHAPTER 6 MEMS FUNCTIONAL OCT IMAGING (PS, DOPPLER)

In this chapter, implementation of functional OCT, including PS OCT and Doppler

OCT will be introduced. The goal of this study is to explore the potential of our existing

time domain OCT system, to improve the image contrast for special tissues, such as the

ones with birefringent properties, and blood vessels. To realize PS OCT, several optical

components need to be added to the existing OCT configuration. While the major

difference for implementing Doppler OCT from the basic system is signal processing

algorithm. In this study, ex vivo PS OCT animal imaging using MEMS probe has been

realized; while Doppler OCT imaging results are obtained by using translational stage

for lateral scans.

PS OCT System

PS OCT System Configuration

In order to add polarization sensitive functionality to the existing system, several

optical components are added to the system. These components are one polarizer, one

polarization modulator, three sets of polarization controllers, two polarization beam

splitters and another heterodyne photodetector. The complete PS OCT system is

demonstrated in Figure 6-1. Stokes vectors are employed to represent the birefringent

properties of the sample. Using Stokes vectors to describe light polarization is an

alternative way to Jones matrix. For a polarization beam

y

x

i

y

i

x

eE

eE

, its Stokes vectors are

defined as

115

I Ex2 Ey

2

Q Ex2 Ey

2

U 2ExEy cos( x y )

V 2ExEy sin( x y )

(6-1)

2222 VUQI (6-2)

Three of the four parameters are independent, and could be used to derive the

fourth one. The results of the PS OCT system are two sets of Stokes vectors images,

with linearly polarized input light and circularly polarized input light.

Figure 6-1. Schematic of PS OCT system.P: polarizer; OC: optical circulator; PC: polarization controller; BS: 2 x 2 beam splitter; RSOD: rapid scanning optical delay line; PBS: polarization beam splitter; DA and DB: balanced photodetectors.

Infrared light coming from the light source goes through a polarizer first, in order to

obtain linearly polarized light. A polarization modulator is right after it. The polarization

axis of the modulator has a 45° angle to the polarization axis of the input light. When no

voltage is applied on the modulator, input light is not modulated; linearly polarized light

goes through the system. When a voltage is applied to the modulator, a phase shift

116

can be induced, and polarization state will change from linear to circular. The

interference signals from these two polarization states are used to calculate the Stokes

vectors. The polarization controllers in the system are used for tuning so that the

reference signal detected by D1, D2 and D3, D4 are equal. Jones matrix can be used to

represent the polarization states at each stage, when linear polarization is used. The

input light is

00

00

E

E, the reference signal reaches PBS1 can be written as

0

0

2

10

~

~

E

eE

E

E i

r

r

, and the reference signal at PBS2 can be written as

i

i

r

r

eE

eE

E

E

0

0

4

30

~

~

; similarly, the sample signal at PBS1 is

2

1

2

1

2

1

~

~

i

i

s

s

eE

eE

E

E, and the

sample signal at PBS2 is

)(

2

1

4

3

2

1

~

~

i

i

s

s

eE

eE

E

E. The reference and sample signals with the

same polarization orientation will interfere with each other, signals reach the

photodetectors from D1, D2, D3 and D4 are shown in equation (6-3) and equation (6-4),

and then heterodyne detection occur at DA and DB, the DC signal is cancelled at this

point, and the detected signals are shown in equation (6-5).

)cos(2

)cos(2

220

2

2

2

02

0110

2

1

2

01

kzEEEED

kzEEEED

(6-3)

)cos(2

)cos(2

220

2

2

2

04

0110

2

1

2

03

kzEEEED

kzEEEED (6-4)

)cos(4

)cos(4

22042

011031

kzEEDDD

kzEEDDD

B

A

(6-5)

117

Signal processing using Hilbert transform gives the complex signal as shown

below.

)](exp[4~

)](exp[4~

220

0110

kziEED

kziEED

B

A

(6-6)

With these results, a set of stokes parameters of linearly polarization can be

obtained using the following equation.

)sin(16]~~

Im[2

)cos(16]~~

Re[2

)(16~~~~

)(16~~~~

20121

2

0

*

20121

2

0

*

2

2

2

1

2

0

**

2

2

2

1

2

0

**

EEEDDV

EEEDDU

EEEDDDDQ

EEEDDDDI

BA

BA

BBAA

BBAA

(6-7)

During the next time period, voltage is applied on the polarization modulator to

induce circular polarization. A prime is used to distinguish circular polarization from

linear polarization. The input circularly polarized light can be represented by

00

200

E

eEi

.

Similar to the linear case, the reference light reaches PBS1 is

0

2

0

2

10

~

~

E

eE

E

E i

r

r

, and the

reference light at PBS2 is

i

i

r

r

eE

eE

E

E

0

2

0

4

30

~

~. The sample signals at PBS1 and PBS2 are

2

1

2

1

2

1

~

~

i

i

s

s

eE

eE

E

E and

2

1

2

1

4

3

~

~

i

i

s

s

eE

eE

E

E. Again, the same polarization orientation will

interfere with each other, and the interference signals are

)cos(2

)2

cos(2

220

2

2

2

02

0110

2

1

2

01

kzEEEED

kzEEEED

(6-8)

118

)cos(2

)2

cos(2

220

2

2

2

04

0110

2

1

2

03

kzEEEED

kzEEEED

(6-9)

The detected signals from circularly polarized input light are

)cos(4

)2

cos(4

22042

011031

kzEEDDD

kzEEDDD

B

A

(6-10)

The complex signal after Hilbert transform can be express as

)](exp[4~

)]2

(exp[4~

220

0110

kziEED

kziEED

B

A

(6-11)

The signal Stokes vectors with circular polarization input light can be calculated by

the following equation.

)2

sin(16]~~

Im[2

)2

cos(16]~~

Re[2

)(16~~~~

)(16~~~~

20121

2

0

*

20121

2

0

*

2

2

2

1

2

0

**

2

2

2

1

2

0

**

EEEDDV

EEEDDU

EEEDDDDQ

EEEDDDDI

BA

BA

BBAA

BBAA

(6-12)

When both the linearly polarized signal and the circularly polarized signal are

collected, two sets of Stokes vectors, 8 images in total are displayed on the computer

screen. With 4 images in the first row representing the Stokes parameters: I, Q, U, V

obtained with linearly polarized light, and another 4 images in the second row

representing those obtained with circularly polarized light.

Ex Vivo PS OCT Imaging Experiment

Ex vivo 2D and 3D PS OCT imaging of canine meniscus have been performed to

show the capability of the system. During the imaging session, a canine knee joint was

119

cut open (Figure 6-2) to expose the menisci. The 2nd generation OCT probe was

employed for this experiment. It was positioned at different locations of the menisci for

imaging.

Figure 6-2. PS OCT imaging of canine meniscus.

Figure 6-3. PS OCT imaging results. A) Stokes parameters of linear and circular polarization, B) 3D reconstruction of intensity image, and C) 3D reconstruction of parameter Q.

A

B

C

I Q U V

120

A lateral scanning area of 2.3 × 2.3 mm2 was obtained by driving two opposite

pairs of the actuators with two sets of voltages, i.e., 0~4 V for the circumferential

direction scanning and 0.5~3.5 V for the longitudinal direction scanning. Figure 6-3

shows 2D and 3D PS OCT images of the canine meniscus. The 2D Stokes parameter

images, corresponding to I, Q, U and V, respectively, are shown in Figure 6-3 A, where

each row represents one of the two polarization states. Figure 6-3 B shows the 3D OCT

reconstructed from the intensity parameter I, and Figure 6-3 C shows the 3D

reconstruction from parameter Q. The two shadowed areas in Figure 6-3 A are artifacts

due to two blots on the FEP tubing.

Doppler OCT System

Doppler OCT Principle

To realize Doppler OCT, no hardware modification is required for the original time

domain OCT system. Only an additional signal processing step to obtain Doppler

information is required. In our Doppler OCT system, phase resolved signal processing

method is adopted. The principle of this algorithm has been discussed in the section

Functional OCT: Doppler OCT. Doppler frequency could be calculated from the

equation: Tf KKD /1 . Hilbert transform is used to obtain the phase information

, once signal )(t is collected by the data acquisition card. During signal processing,

Hilbert transform is realized by fast Fourier transform (FFT). It could be done according

to the following procedure:

)(~

)()( 1 tFFTlterBandPassFiHFFTt

In this method, Hilbert transform is realized in the frequency domain. First, a

Fourier transform is conducted, and then the signal is multiplied with the function

121

01

00

H

. A band pass filter is added to reduce noise. Finally, an inverse Fourier

transform provides the complex signal with its phase information. Each four A-lines

were used for averaging in image display.

In Vivo Doppler OCT Imaging Experiment

In vivo experiments on imaging mouse blood vessels were conducted in our lab.

Transverse scans were realized by transverse scanning PI stage. The sample was a

female athymic (nu/nu) nude mouse with a body weight of 20 g to 26 g (Harlan

Laboratories, Indianapolis, IN). The mouse was anesthetized by injecting ketamine (100

mg/kg) and xylazine (10 mg/kg) intraperitoneally, shown in Figure 6-4 A. Imaging results

of Doppler OCT were shown in Figure 6-4 B and C; image depth is 1.6 mm and image

width is 2.3 mm. From the comparison between the intensity image and the Doppler

image, we can see that the contrast of blood vessel from the Doppler image has been

greatly improved.

Figure 6-4. In vivo Doppler imaging. A) sample mouse during imaging session, B) intensity image through the window chamber, and C) Doppler image through the window chamber.

Blood Vessels

A C

B

122

Summary

In this chapter, detailed implementations of PS OCT and Doppler OCT at the BML

were introduced. PS OCT was successfully demonstrated with the 2nd generation

MEMS probe. 2D and 3D OCT images of canine meniscus were presented. Improved

contrast of Doppler OCT for rat blood vessels was also successfully demonstrated.

123

CHAPTER 7 MEMS OCT/PHOTOACOUSTIC IMAGING

So far we have discussed OCT’s ability for high resolution endoscopic imaging.

For highly scattering tissues, the penetration depth of OCT could be limited. It is

therefore desirable to use another imaging modality with higher penetration depth to

work with OCT as an extension for deeper tissue imaging. Photoacoustic (PA) imaging

is a suitable candidate for this role. PA imaging is a combination of optical imaging and

ultrasound imaging. It can normally reach 3 – 5 mm deep in tissue, with about 100 µm

resolution. With the help of PA imaging, deeper tissue structures that are not visible with

OCT could be seen, and it could help to make the decision of whether more tests need

to be conducted for deeper tissues. In this chapter, we introduce a prototype

endoscopic probe that could be used for both OCT and PA imaging. In the following

sections, first the background and principle of PA imaging will be introduced. Then the

photoacoustic imaging system employed in this study will be presented, both ex vivo

and in vivo experimental results will be shown to demonstrate the system capability.

And the OCT/PA imaging probe will be introduced at last.

Photoacoustic Imaging System

Principle and Background

PA imaging is based on the PA effect, which describes the generation of acoustic

waves by the incidence of optical energy. In the case of PA imaging, when laser output

is incident on to the tissue, the absorption of the optical energy causes the tissue to

have thermal expansion, and subsequently, emitting broadband ultrasonic waves. The

time of flight, amplitude etc. of the ultrasonic waves provide information regarding

124

location, strength and dimension of the acoustic source, therefore tissue reconstruction

can be realized by detecting the ultrasound signals[162]-[165].

PA imaging combines high optical absorption contrast and high ultrasonic spatial

resolution, and therefore has received a lot of attention recently. Researchers have

demonstrated PA imaging applications in breast cancer, brain vasculature, skin disorder

and arthritis, etc. [164]-[171]. For endoscopic photoacoustic imaging, methods such as

using a geared micromotor to scan a single transducer and using a galvanometer for

lateral scan have been proposed. However, using micromotor can only realize 2D side-

view scan unless an external linear stage is used; and using galvanometer limits the

further miniaturization of the probe due to its relatively large size. Therefore, MEMS

mirror become the ideal choice for this task.

MEMS Based Photoacoustic Probe

An endoscopic probe based on the 3rd generation LSF-LVD MEMS mirror is

designed and assembled for photoacoustic imaging. This is a front view design, as

shown in Figure 7-1. The inner diameter of the probe is 9.5 mm and the outer diameter

is 11.5 mm. A ring-shaped polyvinylidene fluoride (PVDF) transducer is attached to the

edge of the probe; it is a one-element ultrasound detector, which is fabricated with a

110 μm-thick Ag ink printed PVDF film (Measurement Specialties Inc.) as the sensing

unit. Flexible PCB of the same shape is attached to it as positive and negative electrode

connectors and backing materials. The central frequency of this detector is 2.5 MHz.

When light comes in, it first incidents on a fixed mirror, then the light is reflected to the

MEMS mirrors for transverse scans. Ultrasound signal generated from the tissue due to

thermal expansion is detected by the PVDF transducer. During scanning, the four

actuators are driven by 0 ~ 4 V differential ramp voltages and optical scan angles of ±

125

31° could be achieved. When the imaging surface is place 9 mm away from the MEMS

mirror, scanning area of up to 9 × 9 mm2 could be obtained. The scanning beam size is

1 × 1 mm2, determined by the MEMS mirror aperture size.

Figure 7-1. Endoscopic photoacoustic probe. A) schematic of probe, B) 3D model, and C) assembled probe.

Photoacoustic System

The endoscopic photoacoustic system employed in this study is demonstrated in

Figure 7-2. The laser employed was a Nd:YAG 532 nm laser (NL 303HT from EKSPLA,

Lithuania); it generated pulses of 20 ns duration at 10 Hz. Two function generators

(AFG3022B, Tektronics) were used to generate synchronization signals for the data

acquisition and driving signal for MEMS mirror. The ultrasound signal was first amplified

by a pre-amplifier with a gain of 17 dB and then further amplified by an amplifier with a

controllable gain from 5 dB to 20 dB. Sampling rate of the data acquisition card was 50

MHz; and a band-pass filter (Panametrics, Waltham, Massachusetts) from 1 MHz to 5

MHz was used to pass the signal and reduce noises.

B

MEMS mirror

Fixed mirror

Glass cover

PCB

Transducer

MEMS mirror

Tissue A C

126

Figure 7-2. Photoacoustic system.

Photoacoustic Experiments

A series of ex vivo and in vivo experiments were conducted in the Biomedical

Optics Laboratory (directed by Dr. Huabei Jiang) to demonstrate the system capability.

In these experiments, the incident laser beam was 1 mm × 1 mm in size, and the

incident power on the sample was 15 mJ/cm2, which is below the American National

Standards Institute safety limit of 20 mJ/cm2. For 2D images, 50 points of signal were

obtained in each direction; the scanning frequency for fast scan was 100 mHz, and for

slow scan, it was 2 mHz. The scanning time was limited to 250 s, due to the 10 Hz

repetition rate of the laser. During the experiments, ultrasound gel was applied between

the probe and the sample to minimize ultrasound attenuation.

To demonstrate the system 3D capability, an experiment was conducted with a

pencil lead (0.7 mm in diameter) embedded inside a piece of chicken at 1.5 mm

underneath the surface. Figure 7-3 shows the imaging results, which accurately

127

reflected the embedded pencil lead; the reconstructed pencil diameter was ~ 0.72 mm.

In vivo experiment was also conducted, shown in Figure 7-4. The reconstructed 3D

image clearly reflected the size and shape of the blood vessel labeled in the picture.

Figure 7-3. Ex vivo photoacoustic experiment with pencil lead embedded in a piece of chicken. A) photograph of the imaging target, B)-D) coronal, sagittal and cross-section views of the recovered 3D image, and E) 3D rendering of the recovered image.

Figure 7-4. In vivo photoacoustic experiment. A) imaging target, B) reconstructed blood vessel image.

MEMS Probe for Both OCT and Photoacoustic Imaging

MEMS Probe Design

A side-view endoscopic probe was designed for this study. Components for two

imaging modalities were combined into one design, including single mode fiber for OCT

imaging, multi mode fiber and transducer for PA imaging, and GRIN lens and MEMS

mirror for both uses. The two fibers were bundled together and connected to the GRIN

lens inside the probe; and outside, they were connected to the OCT system and the PA

A

B

C

D

E

A B

reconstructed blood vessel

128

system separately. For OCT imaging, this probe works in a similar way as the

previously reported side-view probes, the back-scattered optical signal is collected

through the GRIN lens back into the single mode fiber; for PA imaging, ultrasound

signal is detected by the transducer around the fibers. The MEMS mirror employed for

this probe is the 4th generation MEMS mirror; because for PA imaging, some ultrasound

gel is needed to prevent signal attenuation, and the MEMS mirror should be protected

from the gel. The 4th generation MEMS mirror has its mirror plate hidden under the

frame, and can be easily protected by a covering glass slide.

Figure 7-5. Probe for OCT and PA imaging. A) schematic of the probe, B) pictures of the assembled probe.

The diameter of the probe is about 7 mm. And the designed working distance from

the MEMS mirror to the imaging sample is ~ 3.5 mm. With the MEMS mirror scanning

±15° optical angle, the imaging area is about 1.8 mm × 1.8 mm. Due to the assembly

variation, the working distance for the assembled probe is only about 2 mm. And the

images obtained with this probe are about 1mm × 1mm.

Ex Vivo Experiment

The assembled probe was employed in ex vivo experiments to demonstrate its

capability for both OCT and PA imaging. The same procedures as that described in

A

B

MEMS mirror

transducer

GRIN lens

GRIN lens

transducer

129

Chapter 4 for the 4th generation probe was adopted for this experiment as the same

type of MEMS mirror was employed. From Figure 7-6, layered structure could be clearly

seen in the 2D image. The image looks stretched due to the smaller imaging range. The

PA imaging experiments are on-going at the Biomedical Optics Laboratory, and results

will be updated in future publications.

Figure 7-6. OCT images of mouse ear. A) 2D OCT image of mouse ear, B) 3D OCT image of mouse ear. (E: epidermis, D: dermis, cc: conjunction cartilage, C: cartilage).

Summary

This chapter introduced the concept of combining OCT and PA imaging with one

endoscopic probe. First, PA imaging using MEMS scanner was successfully

demonstrated. And then, a side view endoscopic probe that has the ability for both OCT

and PA imaging was introduced. 2D and 3D OCT images of a mouse ear were obtained

with the prototype MEMS probe.

A

E

D

D E

cc

cc C

200 µm

B

130

CHAPTER 8 CONCLUSION

The major goal of this study is to explore the applications of OCT system

combined with MEMS technology in the areas of endoscopic imaging and refractive

index measurement. This chapter highlights the research accomplishments and outlines

the recommended future work.

Summary of Work

Firstly, four generations of electrothermally based MEMS probes have been

developed. The probe size has been reduced from 5 mm to 2.5 mm in diameter. And

the packaging scheme has been simplified over time. Wire-bonding process has been

eliminated for the 4th generation probe, and electrical and mechanical connections have

been made easy with a flip-chip bonding mirror design. The optical performance of the

probes has been analyzed in detail. And the capabilities of the miniature probes for

endoscopic imaging have been successfully demonstrated with in vivo imaging

experiments; layered structures from sample tissues were observed, and tumor regions

were clearly separated from healthy tissue structures.

Secondly, MEMS mirror based free space OCT for refractive index (RI)

measurement in brain tissue slices has been demonstrated. RIs of both white matter

and gray matter were measured under free-loading and compressed status. This work

provides information that can be used to correct distorted optical images caused by

tissue heterogeneity or compression from surgery or certain diseases. Information on RI

change under compression can also help the mechanical property tests of soft tissues

based on elastography.

131

Thirdly, two types of functional OCT were demonstrated. They are polarization

sensitive OCT and Doppler OCT. The implementations of these two types of OCT

systems help improved the image contrast. PS OCT is especially useful for birefringent

tissues, while Doppler OCT helps to detect blood vessels with a higher contrast.

Lastly, OCT was combined with PA imaging to integrate two types of imaging

modalities in one probe. By combining two imaging modalities, we can take advantage

of the high resolution of OCT imaging, and the good imaging depth of PA imaging. The

successful implementations of ex vivo imaging experiments demonstrated the feasibility

of this idea.

Future Work

The successful demonstration of endoscopic OCT imaging using eletrothermally

actuated MEMS mirror opens up new opportunities for early cancer detection and

imaging guided surgery. Despite the achievements of this work, for MEMS OCT probes

to be widely used clinically, more work needs to done.

Firstly, more robust assembly procedures are needed to improve the consistency

in imaging qualities. The current assembly process is susceptible to manual variations.

For example, by using a fiber-GRIN lens module instead of two separate components

would help reduce the effect of focal length change due to the changing distance

between the two components.

Secondly, a different probe housing configuration needs to be developed to reduce

astigmatism introduced by the cylindrical shaped tubing. Using flat tubing and adding

astigmatism correction lenses are two potential resolutions that could be explored.

132

Thirdly, an imaging rendering algorithm that takes into account the non-telecentric

scan effect and non-linearity of the MEMS scan should be developed to reflect the real

size and shape of the sample being imaged.

Fourthly, a front view imaging probe should be developed for the PS OCT

research. A side view probe was used to image canine meniscus ex vivo in this study. A

front view probe is more suitable for imaging the meniscus in live animals.

Fifthly, in the refractive index measurement study, only coronally sectioned brain

tissue slices were measured. Since brain tissues have variously aligned micro-

structures, the effect of the directionality of brain tissue slices on RI and RI changes in

response to compression should be studied in the future.

Lastly, the first prototype probe for OCT and PA imaging has a relatively large

diameter. For clinical use, new designs with smaller sizes are desirable.

133

APPENDIX PUBLICATIONS

The following is a list of archival journal and conference publications generated

from the accomplished research effort.

Journal papers:

1. J. Sun, S. J. Lee, L. Wu, M. Sarntinoranont, and H. Xie, "Refractive index measurement of acute rat brain tissue slices using optical coherence tomography," Opt. Express 20, 1084-1095 (2012)

2. J. Sun, and H. Xie, "MEMS-based endoscopic optical coherence tomography", Int. J. Opt. 2011, 825629, (2011).

3. J. Sun, S. Guo, L. Wu, L. Liu, S. Choe, B. Sorg, and H. Xie, "3D in vivo optical coherence tomography based on a low-voltage, large-scan-range 2D MEMS mirror," Opt. Express 18, 12065-12075 (2010).

4. S. J. Lee, J. Sun, J. Flint, S. Guo, H. Xie, M. King, and M. Sarntinoranont, "Optically based-indentation technique for acute rat brain tissue slices and thin biomaterials", J. Biomed. Mater. Res. B, 97(1), (2011).

5. L. Xi, J. Sun, Y. Zhu, L. Wu, H. Xie, and H. Jiang, "Photoacoustic imaging based on MEMS mirror scanning," Biomed. Opt. Express 1, 1278-1283 (2010).

6. L. Liu, L. Wu, J. Sun, E. Lin and H. Xie, "Miniature endoscopic optical coherence tomography probe employing a two-axis microelectromechanical scanning mirror with through-silicon vias", J. Biomed. Opt. 16, 026006 (Feb 01, 2011).

Conference papers and abstracts:

1. J. Sun, C. Duan, S. Salmuelson, and H. Xie, “Design and Characterization of MEMS Based Optical Coherence Tomography Endoscopic Probe”, to be presented at Biomedical optics (BIOMED 2012), Miami, FL, April 28-May 2, 2012.

2. J. Sun, S. Guo, L. Wu, Choe, B. Sorg and H. Xie, "In vivo 3D and Doppler OCT imaging using electrothermal MEMS scanning mirrors", Proc. SPIE 7594, 759405 (2010).

3. J. Sun, S. Lee, M. Sarntinoranont, and H. Xie, "In vitro refractive index measurement of acute rat brains using optical coherence tomography", presented at the 2010 Biomedical Engineering Society Annual Meeting, Austin, TX, 6-9, Oct. 2010.

4. J. Sun, L. Wu, and H. Xie, "In vivo 3D and Doppler OCT imaging using electrothermal MEMS scanning mirrors", Proc. SPIE 7594, 759405 (2010).

134

5. S. Guo, J. Sun, A. Pozzi, H. Ling, L. Wu, L. Liu, and H. Xie, "3D polarization-sensitive optical coherence tomography of canine meniscus based on a 2D high-fill-factor microelectromechanical Mirror," Engineering in Medicine and Biology Society, 2009. (EMBC 2009. Annual International Conference of the IEEE, 3-6 Sept. 2009), pp.1445-1448.

6. S. J. Lee, J. Sun, M. King, H. Xie, and M. Sarntinoranont, "Viscoelastic property changes of acute rat brain tissue slices as a function of cell viability", presented at ASME 2011 Summer Bioengineering Conference (SBC2011), Famington, Pennsylvania, USA, 22-25 June, 2011.

7. S. Lee, J. Sun, H. Xie, and M. Sarntinoranont, "An optically-based indentation technique for thin soft biomaterials", presented at 2009 Biomedical Engineering Society Annual Fall Meeting, Pittsburgh, PA, 2009.

8. S. Guo, L. Wu, J. Sun, L. Liu, and H. Xie, "Three-dimensional optical coherence tomography based on a high-fill-factor microelectromechanical mirror”, in Novel Techniques in Microscopy, OSA Technical Digest (CD) (Optical Society of America, 2009), paper NTuB3.

9. L. Wu, S. Samuelson, J. Sun, K. Jia, W. Lau, S. Choe, B. Sorg, and H. Xie, “A 2.8mm imaging probe based on a high-fill-factor MEMS mirror and wire-bonding-free packaging for endoscopic optical coherence tomography”, Micro Electro Mechanical Systems (MEMS), 2011 IEEE 24th International Conference on (Cancun, Mexico, 23-27 Jan. 2011), pp.33-36.

10. D. Wang, L. Fu, J. Sun, H. Jia, and H. Xie, "Design optimization and implementation of a miniature optical coherence tomography probe based on a MEMS mirror", Proc. SPIE 8191, 81910M (2011).

11. W. Liao, W. Liu, L. Xi, J. Sun, Y. Zhu, L. Wu, H. Jiang, and H. Xie, "Miniature photoacoustic imaging probe using MEMS scanning micromirror", Optical MEMS and Nanophotonics (OMN), 2011 International Conference on (Istanbul, Turkey, 8-11 Aug. 2011), pp.115-116.

135

LIST OF REFERENCES

[1] "Cancer facts and figures 2012,” Atlanta: American Cancer Society; 2012.

[2] D. Huang, E. A. Swanson, C. P. Lin, J. S. Schuman, W. G. Stinson, W. Chang, M. R. Hee, T. Flotte, K. Gregory, C. A. Puliafito, and J. G. Fujimoto, "Optical coherence tomography", Science 254, 1178-1181 (1991).

[3] P. Rai-Choudhury, "MEMS and MOEMS: Technology and Applications," SPIE Press, 2000.

[4] J. M. Bustillo, R. T. Howe, and R. S. Muller, "Surface micromachining for microelectromechanical systems," Proceedings of the IEEE 86, 1552-1574, (1998).

[5] G. T. A. Kovacs, N. I. Maluf, and K. E. Petersen, "Bulk micromachining of silicon," Proceedings of the IEEE, 86, 1536-1551, (1998).

[6] F. Koenig, J. Knittel, and H. Stepp, "Diagnosing cancer in vivo," Science 292, 1401-1403, (2001).

[7] V. Backman, L. T. Perelman, J. T. Arendt, R. Gurjar, M. G. Muller, Q. Zhang, G. Zonios, E. Kline, T. McGillican, T. Valdez, J. V. Dam, M. Wallace, K. Badizadegan, J. M. Crawford, M. Fitzmaurice, S. Kabani, H. S. Levin, M. Seiler, R.R.Dasari, I. Itzkan, and M. S. Feld, "Detection of preinvasive cancer cells," Nat. Biotechnol 406, 35-36, (2000).

[8] N. Claxton, T. Fellers, and M. Davidson, "Laser Scanning Confocal microscopy." Olympus, (2006).

[9] A. Boyde, L. Lovicar and J. Zamecnik, "Combining confocal and BSE SEM imaging for bone block surfaces, Eur Cells Mater 9, 33–38 (2005).

[10] Z. Sbeity, P. M. Palmiero, C. Tello, J. M. Liebmann, and R. Ritch, "Non-contact in vivo confocal scanning laser microscopy in exfoliation syndrome, exfoliation syndrome suspect and normal eyes." Acta Ophthalmol 89(3), 241, (2011).

[11] T. Wilson and A. R. Carlini, "Three dimensional imaging in confocal imaging systems with finite-sized detectors," J. Microscopy, 141, 51-66, (1988).

[12] M. Gu, X. Gan, A. Kisteman, M. G. Xu, "Comparison of penetration depth between two-photon excitation and single-photon excitation in imaging through turbid tissue media," Appl. Phys. Lett. 77, 1551-1553, (2000).

[13] F. Helmchen, and W. Denk, "Deep tissue two-photon microscopy," Nat. Methods 2, 932-940 (2005).

136

[14] P. J. Campagnola, M.Wei, A. Lewis, L. M. Loew, "High-Resolution Nonlinear Optical Imaging of Live Cells by Second Harmonic Generation," Biophys. J. 77(6), 3341-3349, (1999).

[15] X.Deng, E. D. Williams, E. W. Thompson, and X. G. M. Gu, "Second-harmonic generation from biological tissues: Effect of excitation wavelength," Scanning 24, 175-178, (2002).

[16] W. R. Zipfel, R. M. Williams, R. Christie, A. Y. Nikitin, B. T. Hyman, and W. W. Webb, "Live tissue intrinsic emission microscopy using multiphoton-excited native fluorescence and second harmonic generation," Proc. Natl. Acad. Sci. U.S.A. 100, 7075–7080 (2003).

[17] J. Cheng and, X. S. Xie, " Coherent Anti-Stokes Raman Scattering Microscopy:  Instrumentation, Theory, and Applications", J. Phys. Chem. B 108(3), 827-840, (2004).

[18] E. O. Potma, C. L. Evans, and X. S. Xie, "Heterodyne coherent anti-Stokes Raman scattering (CARS) imaging," Opt. Lett. 31, 241-243 (2006).

[19] A. Puliafito, M. R. Hee, C. P. Lin, E. Reichel, J. S. Schuman, J. S. Duker, J. A. Izatt, E. A. Swanson, and J. G. Fujimoto, "Imaging of macular diseases with optical coherence tomography," Ophthalmology 102, 217-229 (1995).

[20] J. S. Schuman, M. R. Hee, A. V. Arya, T. Pedut-Kloizman, C. A. Puliafito, J. G. Fujimoto, and E. A. Swanson, "Optical coherence tomography: a new tool for glaucoma diagnosis," Curr. Opin. Ophthalmol. 6, 89-95 (1995).

[21] J. G. Fujimoto, "Optical coherence tomography for ultrahigh resolution in vivo imaging," Nat. Biotechnol. 21, 1361-1367 (2003).

[22] E. A. Swanson, D. Huang, M. R. Hee, J. G. Fujimoto, C. P. Lin, and C. A. Puliafito, "High-speed optical coherence domain reflectometry," Opt. Lett. 17, 151-153 (1992).

[23] J. Welzel, "Optical coherence tomography in dermatology: a review," Skin Res. Technol. 7, 1- 9 (2001).

[24] B. H. Park, C. Saxer, S. M. Srinivas, and J. F. de Boer, "In vivo burn depth determination by high-speed fiber-based polarization sensitive optical coherence tomography," J. Biomed. Opt. 6, 474-479 (2001).

[25] M. C. Pierce, R. L. Sheridan, B. H. Park, B. Cense, and J. F. de Boer, "Collagen denaturation can be quantified in burned human skin using polarization-sensitive optical coherence tomography," Burns 30, 511-517 (2004).

[26] M. V. Sivak, Jr., K. Kobayashi, J. A. Izatt, A. M. Rollins, R. Ung-Runyawee, A. Chak, R. C. Wong, G. A. Isenberg, and J. Willis, "High-resolution endoscopic

137

imaging of the GI tract using optical coherence tomography," Gastrointest. Endosc. 51, 474-479 (2000).

[27] S. Jackle, N. Gladkova, F. Feldchtein, A. Terentieva, B. Brand, G. Gelikonov, V. Gelikonov, A. Sergeev, A. Fritscher-Ravens, J. Freund, U. Seitz, S. Schroder, and N. Soehendra, "In vivo endoscopic optical coherence tomography of esophagitis, Barrett's esophagus, and adenocarcinoma of the esophagus," Endoscopy 32, 750- 755 (2000).

[28] X. D. Li, S. A. Boppart, J. Van Dam, H. Mashimo, M. Mutinga, W. Drexler, M. Klein, C. Pitris, M. L. Krinsky, M. E. Brezinski, and J. G. Fujimoto, "Optical coherence tomography: advanced technology for the endoscopic imaging of Barrett's esophagus," Endoscopy 32, 921-930 (2000).

[29] J. M. Poneros, S. Brand, B. E. Bouma, G. J. Tearney, C. C. Compton, and N. S. Nishioka, "Diagnosis of specialized intestinal metaplasia by optical coherence tomography," Gastroenterology 120, 7-12 (2001).

[30] G. J. Tearney, M. E. Brezinski, B. E. Bouma, S. A. Boppart, C. Pitris, J. F. Southern, and J. G. Fujimoto, “In vivo endoscopic optical biopsy with optical coherence tomography,” Science 276, 2037-2039 (1997).

[31] H. L. Fu, Y. Leng, M. J. Cobb, K. Hsu, J. H. Hwang and X. Li, "Flexible miniature compound lens design for high-resolution optical coherence tomography balloon imaging catheter", J. Biomed. Opt. 13, 060502 (2008).

[32] T. Xie, D. Mukai, S. Guo, M. Brenner, and Z. Chen, "Fiber-optic-bundle-based optical coherence tomography," Opt. Lett. 30, 1803-1805 (2005).

[33] A. Sergeev, V. Gelikonov, G. Gelikonov, F. Feldchtein, R. Kuranov, N. Gladkova, N. Shakhova, L. Snopova, A. Shakhov, I. Kuznetzova, A. Denisenko, V. Pochinko, Yu Chumakov, and O. Streltzova, “In vivo endoscopic OCT imaging of precancer and cancer states of human mucosa,” Opt. Express 1, 432-440 (1997).

[34] Y. Wang, M. Bachman, G. -P. Li, S. Guo, B. J. F. Wong, and Z. Chen, "Low-voltage polymer-based scanning cantilever for in vivo optical coherence tomography," Opt. Lett. 30, 53-55 (2005).

[35] X. Liu, M. J. Cobb, Y. Chen, M. B. Kimmey, and X. Li, "Rapid-scanning forward-imaging miniature endoscope for real-time optical coherence tomography," Opt. Lett. 29, 1763-1765 (2004).

[36] J. Wu, M. Conry, C. Gu, F. Wang, Z. Yaqoob, and C. Yang, "Paired-angle-rotation scanning optical coherence tomography forward-imaging probe," Opt. Lett. 31, 1265-1267 (2006).

138

[37] L. J. Hornbeck, "Digital light processing and MEMS: An overview," presented at the IEEE/LEOS 1996 Summer Topical Meetings Keystone, CO, Aug. 5-9, (1996).

[38] S.J.Shellhammer, D. P. Goren, T. Pavlidis, "Novel signal-processing techniques in barcode scanning," Robotics & Automation Magazine, IEEE, 6 (1), 57-65, (1999).

[39] T. M. Lippert, "Virtual retinal display, a scanned photon delivery system," LEOS Annual Meeting. 1, 283-284, (1999).

[40] W.O. Davis, R. Sprague, J. Miller, "MEMS-based pico projector display," Optical MEMs and Nanophotonics, 2008 IEEE/LEOS International Conference on, 31-32, (2008).

[41] J. Bernstein, S. Cho, A.T. King, A. Kourepenis, P. Maciel, M. Weinberg, "A micromachined comb-drive tuning fork rate gyroscope," Micro Electro Mechanical Systems, 1993, MEMS '93, Proceedings An Investigation of Micro Structures, Sensors, Actuators, Machines and Systems. IEEE. 143-148, Feb, (1993).

[42] http://endoscopy.md/index.php

[43] B. J. Catalone, “Flexible Endoscopy: Infection Control and Best Practices,” Sept. 18, (2007).

[44] P. Cotton, Advanced Digestive Endoscopy: Practice and Safety (Wiley-Blackwell, 1st edition, 2008).

[45] Y. Pan, H. Xie, and G. K. Fedder, "Endoscopic optical coherence tomography based on a microelectromechanical mirror," Opt. Lett. 26, 1966-1968 (2001).

[46] K. E. Petersen, "Silicon Torsional Scanning Mirror," IBM J. Res. Dev. 24 (5), 631-637, (1980).

[47] L. Fan and M. C. Wu, "Two-Dimensional Optical Scanner with Large Angular Rotation Realized by Self-Assembled Micro-Elevator," IEEE/LEOS Summer Topical Meetings, 107-108, (1999).

[48] G. D. J. Su, H. Toshiyoshi, and M. C. Wu, "Surface-Micromachined 2-D Optical Scanners with High-Performance Single-Crystalline Silicon Micromirrors," Photonics Technology Letters 13, 606-608, (2001).

[49] D. S. Greywall, C.S. Pai, S. H. Oh, C. P. Chang, D. M. Marom, P. A. Busch, R. A. Cirelli, J. A. Taylor, F. P. Klemens, T. W. Sorsch, J. E. Bower, W. Y. C. Lai, and H. T. Soh, "Monolithic Fringe-Field-Activated Crystalline Silicon Tilting-Mirror Devices," J. Microelectromech. S. 12, 702-707, (2003).

139

[50] D. S. Greywall, C.S. Pai, S. H. Oh, C. P. Chang, D. M. Marom, P. A. Busch, R. A. Cirelli, J. A. Taylor, F. P. Klemens, T. W. Sorsch, J. E. Bower, W. Y. C. Lai, and H. T. Soh, "Monolithic Fringe-Field-Activated Crystalline Silicon Tilting-Mirror Devices," J. Microelectromech. S. 12, 702-707, (2003).

[51] Z. Hao, W. Brian, W. Michael, B. Justin, and H. J. A, "A design methodology for a bulk-micromachined two-dimensional electrostatic torsion micromirror," J. Microelectromech. S. 12, 692-701, (2003).

[52] J. C. Chiou and Y. C. Lin, "A Multiple Electrostatic Electrodes Torsion Micromirror Device With Linear Stepping Angle Effect," J. Microelectromech. S. 12, 913-920, (2003).

[53] M. R. Dokmeci, A. Pareek, S. Bakshi, M. Waelti, C. D. Fung, K. H. Heng, and C. H. Mastrangelo, "Two-Axis Single-Crystal Silicon Micromirror Arrays," J. Microelectromech. S. 13, 1006-1017, (2004).

[54] K. Joudrey, G. G. Adams, and N. E. McGruer, "Design, modeling, fabrication and testing of a high aspect ratio electrostatic torsional MEMS micromirror," J. Micromech. Microeng. 16, 2147-2156, (2006).

[55] R. A. Conant, J. T. Nee, K. Y. Lau and R. S. Muller," A flat high-frequency scanning micromirror, " Hilton Head Solid-State Sensor and Actuator Workshop (2000).

[56] H. Schenk, P. Dun, D. Kunze, H. Laher, and H. Kiick, " An Electrostatically Excited 2D Micro-Scanning-Mirror with an In-plane Configuration of the Driving Electrodes," presented at the 13th IEEE International Conference on Micro Electro Mechanical Systems, (2000).

[57] W. Piyawattanametha, P. R. Patterson, D. Hah, H. Toshiyoshi, and M. C. Wu, "Surface and Bulk- Micromachined Two-Dimensional Scanner Driven by Angular Vertical Comb Actuators," J. Microelectromech. S. 14, 1329-1338, (2005).

[58] V. Milanovic, G. A. Matus, and D. T. McCormick, "Gimbal-Less Monolithic Silicon Actuators for Tip-Tilt-Piston Micromirror Applications," Journal of Selected Topics in Quantum Electronics 10, 462-471, (2004).

[59] H. Xie, Y. Pan, and G. K. Fedder, "A CMOS-MEMS Mirror With Curled-Hinge Comb Drives," J. Microelectromech. S. 12, 450-457, (2003).

[60] P. R. Patterson, D. Hah, N. Hung, H. Toshiyoshi, R. Chao, and M. C. Wu, "A Scanning Micromirror with Angular Comb Drive Actuation," presented at the 15th IEEE International Conference on Micro Electro Mechanical Systems, Las Vegas, NV, (2002).

140

[61] W. Jung, D. T. McCormick, J. Zhang, L. Wang, N. C. Tien, and Z. Chen, "Three-dimensional endoscopic optical coherence tomography by use of a two-axis microelectromechanical scanning mirror," Appl. Phys. Lett. 88, 1-3, (2006).

[62] D. T. McCormick, W. Jung, Y.-C. Ahn, Z. Chen, and N. C. Tien, "A Three Dimensional Real-Time MEMS Based Optical Biopsy System For in vivo Clinical Imaging," Transducers, 203-208, (2007).

[63] A. D. Aguirre, P. R. Hertz, Y. Chen, J. G. Fujimoto, W. Piyawattanametha, L. Fan, and M. C. Wu, "Two-axis MEMS Scanning Catheter for Ultrahigh Resolution Three-dimensional and En Face Imaging," Opt. Express 15, 2445-2453 (2007).

[64] R. A. Miller and Y. C. Tai, "Micromachined Electromagnetic Scanning Mirror," Opt. Eng. 36, 1399-1407, (1997).

[65] A. D. Yalcinkaya, H. Urey, and S. Holmstrom, "Nife Plated Biaxail Magentostatic MEMS Scanner," presented at the14th IEEE Intl. Conf. on Solid-State Sensors, Actuators and Microsystems, Lyon, France, 10-14, (2007).

[66] J. W. Judy and R. S. Muller, "Magnetically Actuated, Addressable Microstructures," J. Microelectromech. S. 6, 249-256, (1997).

[67] I. J. Cho, K. S. Yun, H. K. Lee, J. B. Yoon, and E. Yoon, "A Low-Voltage Two-Axis Electromagnetically Actuated Micromirror with Bulk Silicon Mirror Plates and Torsion Bars," presented at the 15th IEEE International Conference on Micro Electro Mechanical Systems, Las Vegas, NV, 20-24, (2002).

[68] J. Bernstein, W. P. Taylor, J. D. Brazzle, C. J. Corcoran, G. Kirkos, J. E. Odhner, A. Pareek, M. Waelti, and M. Zai, "Electromagnetically Actuated Mirror Arrays for Use in 3-D Optical Switching Applications," J. Microelectromech. S. 13, 526-535, (2004).

[69] S. H. Ahn and Y. K. Kim, "Silicon scanning mirror of two DOF with compensation current routing," J. Micromech. Microeng. 14, 1455-1461, (2004).

[70] H. A. Yang, T. L. Tang, S. T. Lee, and W. Fang, "A Novel Coilless Scanning Mirror Using Eddy Current Lorentz Force and Magnetostatic Force," J. Microelectromech. S. 16, 511-520, (2007).

[71] K. H. Kim, B. H. Park, G. N. Maguluri, T. W. Lee, F. J. Rogomentich, M. G. Bancu, B. E. Bouma, J. F. d. Boer, and J. J. Bernstein, "Two-axis magnetically-driven MEMS scanning catheter for endoscopic high-speed optical coherence tomography," Opt. Express, 15, 18130-18140, (2007.

[72] Y. Watanabe, Y. Abe, S. Iwamatsu, S. Kobayashi,Y. Takahashi, T. Sato, "Electromagnetically Driven 2-axis Optical Beam Steering MEMS Mirror and

141

Dependence of Actuation on Magentic Field", IEEJ Transactions on Sensors and Micromachines 130(4), 107-112 (2010).

[73] T. Kawabata, M. Ikeda, H. Goto, M. Matsumoto, and T. Yada, "The 2-Dimensional Micro Scanner Integrated with PZT Thin Film Actuator," presented at the 9th International Conference on Solid State Sensors and Actuators, Chicago, IL, (1997).

[74] H. J. Nam, Y. S. Kim, S. M. Cho, Y. Yee, and J. U. Bu, "Low Voltage PZT Actuated Tilting Micromirror With Hinge Structure," presented at the 2002 IEEE/LEOS International Conference on Optical MEMS, Lugano, Switzerland, Aug. 20-23, (2002).

[75] N. Kikuchi, Y. Haga, M. Maeda, W. Makishi, and M. Esashi, "Piezoelectric 2-D Micro Scanner for Minimally Invasive Therapy Fabricated Using Femtosecond Laser Ablation," presented at the 12th International Conference on Solid State Sensors, Actuators and Microsystems, Boston, MA, Jun. 8-12, (2003).

[76] Y. Yee, J. U. Bu, M. Ha, J. Choi, H. Oh, S. Lee, and H. Nam, "Fabrication and Characterization of a PZT Actuated Micromirror with Two-axis Rotational Motion for Free Space Optics," presented at the 14th IEEE International Conference on Micro Electro Mechanical Systems, Interlaken, Switzerland, Jan. 21-25, (2001).

[77] J. Tsaur, L. Zhang, R. Maeda, and S. Matsumoto, "2D Micro Scanner Actuated by Sol- Gel Derived Double Layered PZT," presented at the15th IEEE International Conference on Micro Electro Mechanical Systems, Las Vegas, NV, Jan. 20-24, (2002).

[78] J. G. Smits, K. Fujimoto, and V. F. Kleptsyn, "Microelectromechanical flexure PZT actuated optical scanner: static and resonance behavior," J. Micromech. Microeng. 15, 1285-1293, (2005).

[79] K. H. Gilchrist, R. P. McNabb, J. A Izatt and S. Grego, "Piezoelectric scanning mirrors for endoscopic optical coherence tomograph", J. Micromech. Microeng. 19, 095012, (2009).

[80] R. Buser, N. F. d. Rooij, H. Tischauser, A. Dommann, and G. Staufert, "Biaxial Scanning Mirror Activated by Bimorph structures for Medical Applications," Sensor. Actuator. A, 31, 29-34, (1992).

[81] J. Bühler, J. Funk, O. Paul, F. P. Steiner, and H. Baltes, "Thermally Actuated CMOS Micromirrors," Sensor. Actuator. A, 46-47, 572-575, (1995).

[82] M. E. Motamedi, S. Park, A. Wang, M. S. Dadkhah, A. P. Andrews, H. O. Marcy, M. Khoshnevisan, and A. E. Chiou, "Development of Micro-Electro-Mechanical Optical Scanner," Opt. Eng. 36, 1346-1353, (1997).

142

[83] M. Ataka, A. Omodaka, N. Takeshima, and H. Fujita, "Fabrication and Operation of Polyimide Bimorph Actuators for a Ciliary Motion System," J. Microelectromech. S. 2, 146-150, (1993).

[84] G. Lin, C. J. Kim, S. Konishi, and H. Fujita, "Design, Fabrication, And Testing of A C shape Actuator," presented at the 8th International Conference on Solid-State Sensors and Actuators, Stockholm, Sweden, Jun. 25-29, (1995).

[85] S. Schweizer, S. Calmes, M. Laudon, and P. Renaud, "Thermally Actuated Microscanner with Large Angle and Low Consumption," Sensor. Actuator. A. 706, 470-477, (1999).

[86] Q. Liu and Q. A. Huang, "Micro-Electro-Mechanical Digital-to-Analog Converter Based on a Novel Bimorph Thermal Actuator," presented at the IEEE Sensors 2002, Orlando, FL, Jun. 12-1, (2002).

[87] M. Sinclair, "A High Frequency Resonant Scanner Using Thermal Actuator," presented at the 15th IEEE International Conference on Micro Electro Mechanical Systems, Las Vegas, NV, Jan. 20-24, (2002).

[88] A. Jain, A. Kopa, Y. Pan, G. K. Fedder, and H. Xie, "A Two-Axis Electrothermal Micromirror for Endoscopic Optical Coherence Tomography," IEEE Journal of Selected Topics in Quantum Electronics 10, 636-642, (2004).

[89] H. Xie, Y. Pan, and G. K. Fedder, "Endoscopic optical coherence tomographic imaging with a CMOS-MEMS micromirror" Sensor. Acutat. A-Phys. 103(1-2), 237-241, (2003).

[90] T. Xie, H. Xie, G. K. Fedder; Y. Pan, "Endoscopic optical coherence tomography with new MEMS mirror," Electron. Lett. 39 (21), 1535- 1536, (2003).

[91] Y. Xu, J. Singh, C. S. Premachandran, A. Khairyanto, K. W. S. Chen, N. Chen, C. J. R. Sheppard, and M. Olivo, "Design and development of a 3D scanning MEMS OCT probe using a novel SiOB package assembly" J. Micromech. Microeng. 18, 125005, (2008 ).

[92] J Singh, J. H. S. Teo, Y. Xu, C. S. Premachandran, N. Chen, R. Kotlanka, M. Olivo, and C. J. R. Sheppard, "A two axes scanning SOI MEMS micromirror for endoscopic bioimaging " J. Micromech. Microeng. 18, 025001, (2008).

[93] X. J. Mu, G. Y. Zhou, H. H. Feng, Y. S. Xu, A. B. Yu, C. W. Tan, K. W. S. Chen, J. Xie, F. S. Chau, "A 3mm endoscopic probe with integrated MEMS micromirror for optical coherence tomography bioimaging," Procedia Engineering 5, Eurosensor XXIV Conference, 681-684, (2010).

[94] M. F. Wang, Y. Xu, C. S. Prem, K. W. Chen, J. Xie, X, Mu, C. W. Tan, A. Yu, H. Feng, "Microfabricated endoscopic probe integrated MEMS micromirror for optical coherence tomography bioimaging," Engineering in Medicine and

143

Biology Society Annual International Conference of the IEEE, 57-60, Aug. 31 2010-Sept. 4, (2010).

[95] X. Mu, W. Sun, H. Feng, A. Yu, K. W. S. Chen, C. Y. Fu, M. Olivo, "MEMS micromirror integrated endoscopic probe for optical coherence tomography bioimaging," Sensor Actuat A-Phys. 168(1), 202-212, (2011)

[96] D. S. Mamedov, S. D. Yakubovich, P. I. Lapin, M. Wojtkowski, and J. G. Fujimoto, "Novel Near-IR Broad-Band Light Sources for Optical Coherence Tomography Based on Superluminescent Diodes," in Optical Coherence Tomography and Coherence Techniques II, W. Drexler, ed., 5861 of Proc. SPIE (Optical Society of America, 2005), paper MA8.

[97] http://www.thorlabs.com/NewGroupPage9.cfm?ObjectGroup_ID=4285

[98] H. Wang, M. W. Jenkins, and A. M. Rollins, “A combined multiple-SLED broadband light source at 1300 nm for high resolution optical coherence tomography,” Opt. Commun. 281(7), 1896–1900 (2008).

[99] http://www.thorlabs.com/newgrouppage9.cfm?objectgroup_id=1760

[100] G. J. Tearney, B. E. Bouma, S. A. Boppart, B. Golubovic, E. A. Swanson, and J. G. Fujimoto, “Rapid acquisition of in vivo biological images by use of optical coherence tomography,” Opt. Lett. 21, 1408–1410 (1996).

[101] J. P. Heritage, A. M. Weiner, and R. N. Thurston, "Picosecond pulse shaping by spectral phase and amplitude manipulation," Opt. Lett. 10, 609-611 (1985) .

[102] K. F. Kwong, D. Yankelevich, K. C. Chu, J. P. Heritage, and A. Dienes, "400-Hz mechanical scanning optical delay line," Opt. Lett. 18, 558-560 (1993).

[103] A. M. Weiner, D. E. Leaird, J. S. Patel, and J. R. Wullert, "Programmable femtosecond pulse shaping by use of a multielement liquid-crystal phase modulator," Opt. Lett. 15, 326-328 (1990).

[104] G. J. Tearney, M. E. Brezinski, B. E. Bouma, S. A. Boppart, C. Pitris, J. F. Southern, and J. G. Fujimoto, "In vivo endoscopic optical biopsy with optical coherence tomography", Science 27 June 1997: 276 (5321), 2037-2039, (1997).

[105] W. K. Niblack, J. O. Schenk, B. Liu, and M. E. Brezinski, "Dispersion in a grating-based optical delay line for optical coherence tomography," Appl. Opt. 42, 4115-4118 (2003).

[106] T. Xie, Z. Wang, and Y. Pan, "Dispersion compensation in high-speed optical coherence tomography by acousto-optic modulation," Appl. Opt. 44, 4272-4280 (2005).

144

[107] G. J. Tearney, B. E. Bouma, and J. G. Fujimoto, "High-speed phase- and group-delay scanning with a grating-based phase control delay line," Opt. Lett. 22, 1811-1813 (1997).

[108] L. Wang, Z. Ding, G. Shi, and Y. Zhang, "In vivo retinal imaging by optical coherence tomography using an RSOD-based phase modulator", Journal of Zhejiang University - Science A. 10(4), 607-612, (2009).

[109] A. F. Fercher, W. Drexler, C, K. Hitzenberger, and T. Lasser, "Optical coherence tomography - principles and applications" Rep. Prog. Phys. 66, 239, (2003).

[110] R. Huber, M. Wojtkowski, K. Taira, J. Fujimoto, and K. Hsu, "Amplified, frequency swept lasers for frequency domain reflectometry and OCT imaging: design and scaling principles," Opt. Express 13, 3513-3528, (2005).

[111] S. Lee, B. Kim, "Line-Field Optical Coherence Tomography Using Frequency-Sweeping Source," IEEE Journal of Selected Topics in Quantum Electronics 14(1), 50-55, (2008).

[112] H. Lim, M. Mujat, C. Kerbage, E. C. Lee, Y. Chen, T. C. Chen, J.F. de Boer, "High-speed imaging of human retina in vivo with swept-source optical coherence tomography ", Opt Express 14(26), 12902-8, (2006).

[113] Y.Yasuno, S. Makita, T. Endo, G. Aoki, Y. Nakamura, M. Yamanari,M. Itoh, T. Yatagai, "Standard and line-field fourier domain optical coherence tomography," Engineering in Medicine and Biology Society, 2005. IEEE-EMBS 2005. 7224-7226, 17-18 Jan. (2006).

[114] J. Zhang, B. Rao, Z. Chen, "Swept source based fourier domain functional optical coherence tomography," Engineering in Medicine and Biology Society, 2005. IEEE-EMBS 2005. 7230-7233, (2005).

[115] P. A. Keane, H. Ruiz-Garcia, S.R. Sadda, “Clinical applications of long-wavelength (1,000-nm) optical coherence tomography,” Ophthal Surg Las Im 42 (4), 67-74, (2011).

[116] G. Guaqliumi, R. Virmani, “The race to achieve the gold standard in coronary imaging,” Rev Esp Cardiol 62(6), 615-24, (2009).

[117] A. F. Fercher, C. K. Hitzenberger, and S. Y. El-Zaiat, "Measurement of intraocular distances by backscattering spectral interferometry", Opt. Commun. 117, 43 - 48, (1995).

[118] F. Lexer, C. K. Hitzenberger, A. F. Fercher, and M. Kulhavy, "Wavelength-tuning interferometry of intraocular distances," Appl. Opt. 36, 6548-6553, (1997).

145

[119] T. Yoshimura, N. Masuzumi and T. Shigematu, " Imaging a reflecting plate located in scattering media using optical frequency domain reflectometry " Opt. Rev. 4, 227,(1997).

[120] J. F. de Boer, B. Cense, B. H. Park, M. C. Pierce, G. J. Tearney, and B. E. Bouma, "Improved signal-to-noise ratio in spectral-domain compared with time-domain optical coherence tomography," Opt. Lett. 28, 2067-2069 (2003).

[121] R. Leitgeb, C. Hitzenberger, and Adolf Fercher, "Performance of fourier domain vs. time domain optical coherence tomography," Opt. Express 11, 889-894, (2003).

[122] Y. Zhao, Z. Chen, Z. Ding, H. Ren, and J. S. Nelson, "Real-time phase-resolved functional optical coherence tomography by use of optical Hilbert transformation," Opt. Lett. 27, 98-100, (2002).

[123] V. Westphal, S. Yazdanfar, A. M. Rollins, and J. A. Izatt, "Real-time, high velocity-resolution color Doppler optical coherence tomography," Opt. Lett. 27, 34-36, (2002).

[124] Z. Chen, Y. Zhao, S. M. Srinivas, J. S. Nelson, N. Prakash, and R. D. Frostig, "Optical Doppler Tomography", IEEE J. Sel. Top Quant. 5(4), JULY/AUGUST,

(1999).

[125] Z.Chen, T. E. Milner, X. Wang, S. Srinivas, and J. S. Nelson, "Optical Doppler Tomography: Imaging in vivo Blood Flow Dynamics Following Pharmacological Intervention and Photodynamic Therapy," Photochem. Photobiol. 67 (56–60), (1998).

[126] Y. Zhao, Z. Chen, C. Saxer, S. Xiang, J. F. de Boer, and J. S. Nelson, "Phase-resolved optical coherence tomography and optical Doppler tomography for imaging blood flow in human skin with fast scanning speed and high velocity sensitivity," Opt. Lett. 25, 114-116, (2000).

[127] Z. Ding, Y. Zhao, H. Ren, J. Nelson, and Z. Chen, "Real-time phase-resolved optical coherence tomography and optical Doppler tomography," Opt. Express 10, 236-245 (2002).

[128] Y. Zhao, Z. Chen, C. Saxer, Q. Shen, S. Xiang, J. F. de Boer, and J. S. Nelson, "Doppler standard deviation imaging for clinical monitoring of in vivo human skin blood flow," Opt. Lett. 25, 1358-1360 (2000).

[129] M. R. Hee, D. Huang, E. A. Swanson, and J. G. Fujimoto, “Polarization-sensitive low-coherence reflectometer for birefringence characterization and ranging,” J. Opt. Soc. Am. 9, 903–98 (1992).

146

[130] M. Todorović, S. Jiao, L. V. Wang, and G. Stoica, "Determination of local polarization properties of biological samples in the presence of diattenuation by use of Mueller optical coherence tomography," Opt. Lett. 29, 2402-2404 (2004).

[131] J. F. de Boer, T. E. Milner, and J. S. Nelson, "Determination of the depth-resolved Stokes parameters of light backscattered from turbid media by use of polarization-sensitive optical coherence tomography," Opt. Lett. 24, 300-302 (1999).

[132] E. Götzinger, M. Pircher, B. Baumann, C. Ahlers, W. Geitzenauer, U. Schmidt-Erfurth, and C. K. Hitzenberger, "Three-dimensional polarization sensitive OCT imaging and interactive display of the human retina," Opt. Express 17, 4151-4165 (2009).

[133] Y. Wang, C. Lu, M. Tsai, L. Lu, S. Chen, C. Yang, "In vivo Depth-resolved Birefringence of the Hyperlipidemic Rat’s Liver Monitored with Polarization-Sensitive Optical Coherence Tomography," Lasers and Electro-Optics, 2005. CLEO/Pacific Rim 2005. 90- 91, 30-02 Aug. (2005).

[134] C. Hitzenberger, E. Goetzinger, M. Sticker, M. Pircher, and A. Fercher, "Measurement and imaging of birefringence and optic axis orientation by phase resolved polarization sensitive optical coherence tomography," Opt. Express 9, 780-790 (2001).

[135] G. Yao and L.V. Wang, "Two-dimensional depth-resolved Mueller matrix characterization of biological tissue by optical coherence tomography," Opt. Lett. 24, 537-539 (1999).

[136] K. Schoenenberger, B. W. Colston, D. J. Maitland, L. B. Da Silva, and M. J. Everett, "Mapping of Birefringence and Thermal Damage in Tissue by use of Polarization-Sensitive Optical Coherence Tomography," Appl. Opt. 37, 6026-6036 (1998).

[137] J. F. de Boer, T. E. Milner, M. J. C. van Gemert, and J. S. Nelson, "Two-dimensional birefringence imaging in biological tissue by polarization-sensitive optical coherence tomography," Opt. Lett. 22, 934-936 (1997).

[138] H. Ren, Z. Ding, Y. Zhao, J. Miao, J. S. Nelson, and Z. Chen, "Phase-resolved functional optical coherence tomography: simultaneous imaging of in situ tissue structure, blood flow velocity, standard deviation, birefringence, and Stokes vectors in human skin," Opt. Lett. 27, 1702-1704 (2002).

[139] A. Jain, A. Kopa, Y. Pan, G. K. Fedder, and H. Xie, "A Two-Axis Electrothermal Micromirror for Endoscopic Optical Coherence Tomography," IEEE J. Sel. Top. Quant. 10, 636-642, (2004).

147

[140] L. Wu, and H. Xie, “A large vertical displacement electrothermal bimorph microactuator with very small lateral shift,” Sens. Actuators, A 145–146, 371–379 (2008).

[141] K. Jia, S. R. Samuelson and H. Xie, "Single-wafer solution and optical phased array application of micro-mirror arrays with high fill factor and large sub-apertures," in Digest Tech. Papers IEEE MEMS’10 Conference, Hong Kong, Jan 24-28, 180-3, (2010).

[142] M. Motamedi and S. Rastegar, “Light and temperature distribution in laser irradiated tissue: the influence of anisotropic scattering and refractive index,” Appl. Opt. 28, 2230–2237 (1989).

[143] S. R. Arridge and J. C. Hebden, “Optical imaging in medicine: II. Modelling and reconstruction,” Phys. Med. Biol. 42, 841–853 (1997).

[144] H. Dehghani, B. Brooksby, K. Vishwanath, B. W. Pogue, and K. D. Paulsen, " The effects of internal refractive index variation in near-infrared optical tomography: a finite element modelling approach", Phys. Med. Biol. 48 2713, (2003).

[145] F. P. Bolin, L. E. Preuss, R. C. Taylor, and R. J. Ference, "Refractive index of some mammalian tissues using a fiber optic cladding method," Appl. Opt. 28, 2297-2303 (1989)

[146] Q. W. Song, C. Ku, C. Zhang, and R. Michalak, "Modified critical angle method for measuring the refractive index of bio-optical materials and its application to bacteriorhodopsin," J. Opt. Soc. Am. B 12, 797-803 (1995).

[147] H. Li and S. Xie, "Measurement method of the refractive index of biotissue by total internal reflection," Appl. Opt. 35, 1793-1795 (1996).

[148] G. J. Tearney, M. E. Brezinski, J. F. Southern, B. E. Bouma, M. R. Hee, and J. G. Fujimoto, "Determination of the refractive index of highly scattering human tissue by optical coherence tomography," Opt. Lett. 20, 2258-2260 (1995).

[149] J. Binding, J. B. Arous, J. F. Léger, S. Gigan, C. Boccara, and L. Bourdieu, "Brain refractive index measured in vivo with high-NA defocus-corrected full-field OCT and consequences for two-photon microscopy," Opt. Express 19, 4833-4847, (2011).

[150] N. Lue, J. Bewersdorf, M. D. Lessard, K. Badizadegan, R. R. Dasari, M. S. Feld, and G. Popescu, "Tissue refractometry using Hilbert phase microscopy," Opt. Lett. 32, 3522-3524 (2007).

[151] B. Rappaz, P. Marquet, E. Cuche, Y. Emery, C. Depeursinge, and P. Magistretti, "Measurement of the integral refractive index and dynamic cell morphometry of

148

living cells with digital holographic microscopy," Opt. Express 13, 9361-9373 (2005).

[152] J. Beuthan, O. Minet, J. Helfmann, M. Herrig, and G. Müller, “The spatial variation of the refractive index in biological cells,” Phys. Med. Biol. 41(3), 369-382 (1996).

[153] N. L. Dorward, A. Olaf, B. Velani, F. A. Gerritsen,W. F. J. Harkness, N.D. Kitchen, and D. G. T. Thomas, “Postimaging brain distortion: Magnitude, correlates, and impact on neuronavigation,” Journal of Neurosurgery 88, 656–662, (1998).

[154] D. W. Roberts, A. Hartov, F. E. Kennedy, M. I. Miga, and K. D. Paulsen, “Intraoperative brain shift and deformation: A quantitative analysis of cortical displacement in 28 cases,” Neurosurgery, 43(4), 749–758, Oct. (1998).

[155] S. Cheng and L.E. Bilston, “Computational model of the cerebral ventricles in hydrocephalus,” J. Biomed. Eng. 132(5), 054501, May (2010).

[156] Joseph Schmitt, "OCT elastography: imaging microscopic deformation and strain of tissue," Opt. Express 3, 199-211 (1998).

[157] S. J. Lee, J. Sun, J. Flint, S. Guo, H. Xie, M. King, and M. Sarntinoranont, "Optically based-indentation technique for acute rat brain tissue slices and thin biomaterials", J. Biomed. Mater. Res. B, 97(1), (2011).

[158] J. Sun, S. Guo, L. Wu, L. Liu, S. Choe, B. S. Sorg, and H. Xie, "3D In vivo optical coherence tomography based on a low-voltage, large-scan-range 2D MEMS mirror," Opt. Express 18, 12065-12075 (2010).

[159] C. Nicholson, E. Sykova, "Extracellular space structure revealed by diffusion analysis." Trends Neurosci, 21, 207-215, (1998).

[160] T. Shigeno, M. Brock, S. Shigeno, E. Fritschka, and J. Cervós-Navarro, “The determination of brain water content: microgravimetry versus drying-weighing method”, J. Neurosurg. 57(1), 99-107, (1982).

[161] G. Xu, P. Bayly, L. Taber, “Residual stress in the adult mouse brain," Biomech. Model. Mechanobiol. 8(4), 253-62 (2009).

[162] A. A. Oraevsky, S. L. Jacques, R. O. Esenaliev, and F. K. Tittel, “Laser based optoacoustic imaging in biological tissues,” Proc. SPIE 2134A, 122–128 (1994).

[163] A. A. Oraevsky, and A. A. Karabutov, “Optoacoustic tomography,” Chapter 34 in Biomedical Photonics Handbook, T. Vo-Dinh, ed. (CRC Press, Boca Raton, FL, 2003), PM125, 34/1–34/34, (2003).

149

[164] A. A. Oraevsky, A. A. Karabutov, and S. V. Solomatin, “Laser optoacoustic imaging of breast cancer in vivo,” Proc. SPIE 4256, 6–11 (2001).

[165] S. Manohar, A. Kharine, J. C. G. van Hespen, W. Steenbergen, and T. G. van Leeuwen, “The Twente Photoacoustic Mammoscope: system overview and performance,” Phys. Med. Biol. 50(11), 2543–2557 (2005).

[166] S. Manohar, A. Kharine, J. C. G. van Hespen, W. Steenbergen, and T. G. van Leeuwen, “Photoacoustic mammography laboratory prototype: imaging of breast tissue phantoms,” J. Biomed. Opt. 9(6), 1172–1181 (2004).

[167] Q. Zhang, Z. Liu, P. R. Carney, Z. Yuan, H. Chen, S. N. Roper, and H. Jiang, “Non-invasive imaging of epileptic seizures in vivo using photoacoustic tomography,” Phys. Med. Biol. 53(7), 1921–1931 (2008).

[168] Y. Sun, E. Sobel, and H. Jiang, “Quantitative three-dimensional photoacoustic tomography of the finger joints: an in vivo study,” J. Biomed. Opt. 14(6), 064002 (2009).

[169] Y. Sun, and H. Jiang, “Quantitative three-dimensional photoacoustic tomography of the finger joints: phantom studies in a spherical scanning geometry,” Phys. Med. Biol. 54(18), 5457–5467 (2009).

[170] J. A. Viator, G. Au, G. Paltauf, S. L. Jacques, S. A. Prahl, H. Ren, Z. Chen, and J. S. Nelson, “Clinical testing of a photoacoustic probe for port wine stain depth determination,” Lasers Surg. Med. 30(2), 141–148 (2002).

[171] B. Yin, D. Xing, Y. Wang, Y. Zeng, Y. Tan, and Q. Chen, “Fast photoacoustic imaging system based on 320-element linear transducer array,” Phys. Med. Biol. 49(7), 1339–1346 (2004).

[172] E. Z. Zhang, J. G. Laufer, R. B. Pedley, and P. C. Beard, “In vivo high-resolution 3D photoacoustic imaging of superficial vascular anatomy,” Phys. Med. Biol. 54(4), 1035–1046 (2009).

150

BIOGRAPHICAL SKETCH

Jingjing Sun was born in Tianjin, China in January of 1985. She enrolled in Tianjin

University in Tianjin, China in the fall of 2003. Four years later, in July 2007, she

received her Bachelor’s degree in information engineering. After graduation, she joined

the Biophotonics and Microsystems Lab at the University of Florida, and started working

with Dr. Huikai Xie on MEMS based OCT imaging research. In December 2010, she

obtained her Master of Science degree in electrical engineering. And she is expected to

graduate with her Ph.D. degree in electrical engineering in May 2012.