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MEMS Based Optical Coherence Tomography Imaging

Permanent Link: http://ufdc.ufl.edu/UFE0044097/00001

Material Information

Title: MEMS Based Optical Coherence Tomography Imaging
Physical Description: 1 online resource (150 p.)
Language: english
Creator: Sun, Jingjing
Publisher: University of Florida
Place of Publication: Gainesville, Fla.
Publication Date: 2012

Subjects

Subjects / Keywords: coherence -- endoscopy -- index -- mems -- optical -- refractive -- tomography
Electrical and Computer Engineering -- Dissertations, Academic -- UF
Genre: Electrical and Computer Engineering thesis, Ph.D.
bibliography   ( marcgt )
theses   ( marcgt )
government publication (state, provincial, terriorial, dependent)   ( marcgt )
born-digital   ( sobekcm )
Electronic Thesis or Dissertation

Notes

Abstract: Optical coherence tomography (OCT) can provide subsurface cross-sectional information of tissue samples with high resolutions of 1-15 mu. 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 callosum was found to be approximately 4 percent 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 percent strain). For 80 percent strain, measured RIs increased nonlinearly by up to 70 percent and 90 percent 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.
General Note: In the series University of Florida Digital Collections.
General Note: Includes vita.
Bibliography: Includes bibliographical references.
Source of Description: Description based on online resource; title from PDF title page.
Source of Description: This bibliographic record is available under the Creative Commons CC0 public domain dedication. The University of Florida Libraries, as creator of this bibliographic record, has waived all rights to it worldwide under copyright law, including all related and neighboring rights, to the extent allowed by law.
Statement of Responsibility: by Jingjing Sun.
Thesis: Thesis (Ph.D.)--University of Florida, 2012.
Local: Adviser: Xie, Huikai.

Record Information

Source Institution: UFRGP
Rights Management: Applicable rights reserved.
Classification: lcc - LD1780 2012
System ID: UFE0044097:00001

Permanent Link: http://ufdc.ufl.edu/UFE0044097/00001

Material Information

Title: MEMS Based Optical Coherence Tomography Imaging
Physical Description: 1 online resource (150 p.)
Language: english
Creator: Sun, Jingjing
Publisher: University of Florida
Place of Publication: Gainesville, Fla.
Publication Date: 2012

Subjects

Subjects / Keywords: coherence -- endoscopy -- index -- mems -- optical -- refractive -- tomography
Electrical and Computer Engineering -- Dissertations, Academic -- UF
Genre: Electrical and Computer Engineering thesis, Ph.D.
bibliography   ( marcgt )
theses   ( marcgt )
government publication (state, provincial, terriorial, dependent)   ( marcgt )
born-digital   ( sobekcm )
Electronic Thesis or Dissertation

Notes

Abstract: Optical coherence tomography (OCT) can provide subsurface cross-sectional information of tissue samples with high resolutions of 1-15 mu. 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 callosum was found to be approximately 4 percent 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 percent strain). For 80 percent strain, measured RIs increased nonlinearly by up to 70 percent and 90 percent 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.
General Note: In the series University of Florida Digital Collections.
General Note: Includes vita.
Bibliography: Includes bibliographical references.
Source of Description: Description based on online resource; title from PDF title page.
Source of Description: This bibliographic record is available under the Creative Commons CC0 public domain dedication. The University of Florida Libraries, as creator of this bibliographic record, has waived all rights to it worldwide under copyright law, including all related and neighboring rights, to the extent allowed by law.
Statement of Responsibility: by Jingjing Sun.
Thesis: Thesis (Ph.D.)--University of Florida, 2012.
Local: Adviser: Xie, Huikai.

Record Information

Source Institution: UFRGP
Rights Management: Applicable rights reserved.
Classification: lcc - LD1780 2012
System ID: UFE0044097:00001


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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 PHILOSOPH Y UNIVERSITY OF FLORIDA 2012

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2 2012 J ingjing S un

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3 To my parents Xueyi Sun and Xianghui Wang

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4 ACKNOWLEDGMENTS I would like to express my deepest gratitude to my advisor, Dr. Huikai Xie, for his guidance and support through the course o f 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 mem bers, Dr. Henry Zmuda, Dr. Peter Zory, and Dr. Malisa S arntinoranont for their valuable suggestions for my work. I -his well structured lectures provided me a better understanding of the field of optics. I wo uld like to thank Dr. Zory for generously providing his equipments to our lab, a lot of my experiments were conducted with Dr. arntinoranont for her help in the refractive index measurement project; her use ful 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 so lid foundat ion 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. L ee 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

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5 helping me with the animal e xperiments. I would also like to thank Can Duan, who has helped me a lot with experiments since she joined our group. My thank s also go out to everyone in BML and IMG for their useful discussions and friendship s Last but not least, I would like to thank my loving parents and my boyfriend for their love and support.

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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 Las er 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 O utline ................................ ................................ ................................ ......... 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

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7 3 ELECTROTHERMALLY ACTUATED MEMS MIRRORS ................................ ........ 62 Electrothermal Actuation ................................ ................................ ......................... 62 1 st and 2 nd Generations Electrothermal MEMS Mirrors ................................ ........... 63 LSF LVD (3 rd ) and FDSB (4 th ) 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 ................................ ................................ ................. 7 7 In Vivo Experiments ................................ ................................ ................................ 83 Issues with Probe Imaging ................................ ................................ ...................... 86 Effect of the Distance between the Fi ber 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 MEASUREMEN T ................................ ................................ ................................ .. 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

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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

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9 LIST OF TABLES Table page 3 1 Optimized structure parameters of FDSB actuator ................................ ............. 72

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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 schem atic ................................ ................................ ...................... 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 f rom MIT.. ................................ ................................ .. 28 1 9 Electromagnetic MEMS OCT reported by Kim et al [71]. ................................ ... 30 1 10 Electromagnetic MEMS OCT from Yamagata Research I nstitute of Technology, Japan. [72] ................................ ................................ .................... 30 1 11 Piezoelectric MEMS OCT reported by Gilchrist et al [79] ................................ 32 1 12 Electrothermal ME MS 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

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11 2 10 Consecutive A Line scans in Doppler OCT ................................ ........................ 52 2 11 Fr ee 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 1 st generation MEMS mirrors.. ................................ ................................ ............ 64 3 3 2 nd generation: lateral shift free (LSF) MEMS mirror design. .............................. 65 3 4 LSF LVD eletrothermal actua tor 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 1 s t generation probe design. ................................ ................................ ............... 79 4 5 2 nd generation probe design.. ................................ ................................ ............. 80 4 6 3 rd generation probe design.. ................................ ................................ .............. 81 4 7 DFSB probe design.. ................................ ................................ .......................... 82 4 8 DFSB probe design.. ................................ ................................ .......................... 83 4 9 In vivo OCT images of mouse to ngue. ................................ ............................... 84 4 10 In vivo images of mouse ear.. ................................ ................................ ............. 85

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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 test ing experiments. ................................ ................................ .... 89 4 15 Effects of the distance between the fiber and the GRIN lens ............................. 90 4 16 Code V simulation for effect of mir ror 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 sys tem. ................................ ................................ .. 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 Mea surement 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. ................................ .............................. 11 9 6 3 PS OCT imaging results.. ................................ ................................ ................. 119 6 4 In v ivo Doppler imaging. ................................ ................................ ................... 121 7 1 Endoscopic photoacoustic probe. ................................ ................................ ..... 125

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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

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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 J ingjing S un M ay 2012 Chair: Huikai Xie Major: Electrical and Computer Engineering Optical coherence tomography (OCT) can provide subsurface cross sectional information of tissue samples with high resolution s of 1 15 m Microelectromechanical system ( MEMS ) mirrors are small and can perform fast optical scans ; they are suitable for realizing lateral optical scans in OCT systems. In t his thesis MEMS technology is applied to enable endoscopic OCT imaging Two generations of MEMS mirrors have been used. T he 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 tu mor; 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 l ocalized refractive index of acute rat brain tissue slices ; regions including cerebral cort ex, thalamus, corpus callosum, hippocampus and putamen were measured; refractive index in the corpus

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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 8 0% strain). For 80% strain, measured RIs increased nonlinearly by up to 70% and 90% in the cerebral cortex and corpus callosum respectively Further, t wo 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. R at blood vessels were identified by Dopp ler 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 dete ction and real time imaging guided surgery.

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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 5 77 1 90 cancer patients are not expected to live [1] R educ ing cancer death rate has become a global research focus. Early detection and improv ed treatment hav e proven to be effective, as cancer survival rate increased from 49 % to 6 7 % over the past 30 years [1] The development of biomedical optical imaging modalities, especially optical coherence tomography (OCT) ha s been a major contribution to early cancer detection ; a s it can provide in vivo non invasive 3D ultrahigh resolution images that allow s us to see cancer or precancerous lesions that were not visible to traditional imaging techniques, including X ray, c omp uter t omography (CT) scan, ultrasound and m agnetic r esonance i maging (MRI) [2] Endoscopic imaging is very important, because most cancers occur in internal organs; yet it is difficult, because fast 2D lateral s can 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 r esearch motivation and remaining challenges to the existing solutions, then our solut ion will be proposed followed by an outline of this thesis

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17 Traditional C ancer D etection M ethods and T heir L imitations 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 re liable, 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 alternati ves 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 m; and second, the imaging depth should be able to cover the epithelial layer, where most early cancerous morphological changes occur [6] [7] Conventional bio i maging 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. How ever, they are not suitable for early cancer detection due to their low resolution, which is on the order of 100 m. Biomedical optical imaging modalities which can realize ultrahigh resolution imaging (around 10 m) are therefore desirable for this task. Biomedical Optical Imaging C onfocal 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 ultrahig h

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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 m 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 optic al sectioning can be achieve d. Imaging depth of CLSM can reach up to 100 m, which is limited by the scattering effect [8] 3D image can be reconstruct ed 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 ima ges 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 S ample Light Source

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19 Nonlinear Optical Microscopy NLOM uses higher order light matter interaction to generate optical si gnals. Based on no nlinear 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. Compar ed to linear light tissue interaction, the chances of nonlin ear 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 usa ge 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 NL OM. Figure 1 2 A simplified schematic for TPEF and SHG imaging set up Ti : Sapphire Laser Prechirp Unit Ir is filter PMT Bandpass Filter Dichroic Mirror Objective Lens Double clad Photonic Crystal Fiber Scan Mirror Sample

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20 Optical Coherence Tomography OCT employs a n interferometer structure, and relies on coherence gating to reject signal back sc attered 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 tis sue 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 m 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 t he 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, OC T uses relatively low er incident optical power compared to CLSM and nonlinear imaging systems which makes it safer for human tissues imaging Light source 22 Beam splitter Reference mirror Photo detector Lateral scanning DAQ PC

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21 Limitation in OCT Imaging for Endoscopic Application OCT, since its introduction in the early 1990s, has been em ployed extensively in dermatology and ophthalmology [2] [19] [25] It has also been used for imaging internal organs suc h as the GI tracks, bladders, and esophagus [26] [30] [33] [36] As mention ed 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 i s 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 propose d is to use a pair of GRIN lens with proper cut angle at the ends, transverse scanning is achieved by rotating the two lens es at matching speeds [36] Yet, further miniaturization is difficult for this design, b ecause two motors are needed for rotation. In this work, we propose a solution, which integrates MEMS scanners with OCT to achieve fast lateral scan s in small endoscopic probes. Introduction to MEMS Technology M icro electro mechanical systems (MEMS) tech nology makes mechanical or electro mechanical devices structures or system s 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 m any products. For example, MEMS accelerometers can sense changes in velocity, and have been used in automotives for airbag deployment can be adapted to sense the tilting of cell phones, tablets, etc. in o rder to correct the

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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 a re optical scanners. There are two types of optical scanners: digital and analog. Digital Digital Micromirror Devices (DMDs) are digital micromirrors, and they are the core of the Digita l Light Processing (DLP) technology [37] the DLP chips. Analog micromirrors are more flexible, they can steer light continuously over a certain range. Applications including b arcode scanning, virtual retinal display and projection display have been reported [38] [40] Figure 1 4 Existing MEMS devices. A) Draper tuni ng 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 Hert z or A B

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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 engine s MEMS Based Endoscopic OCT Imaging Endoscopic Probes Overv iew 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 size s of endoscopes have different requirements. To look at bod 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 h as 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 g oes 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 prob e s could be designed to be fitted

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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 Co mmercially 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. M EMS 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 bla dder 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. Fig ure 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

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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 re alize 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 (S CS) 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 sca nning 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. Fig ure 1 6 MEMS probe schematic. A) f ront 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 elect rostatic 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 r ational angle of the micromirror reached 2 at resonance and 300V driving A B

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26 voltage. This parallel plate structure has been adopted by many other researchers and rotation angles as large as 8 have been achieved with lower driving voltages in the range of 40V to 200V, but the reported mirror aperture diameters are normally smaller than 500m [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 micromirro rs have been realized using this structure, and the mechanical deflection angle has reached 5.5 at only 16 V resonance [56] and 6.2 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 de flection 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 t he right has a mirror aperture diameter of 800 m. 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 res onance 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,

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27 and injection mold was used for probe body. The diameters of the endosc opic probes range from 3.9 mm to 5.0 mm. 3D in vivo OCT images of healthy rabbit rectal tissue s have been obtained at 8 frames/sec. The image volume size is 1 1 1.4 mm 3 and the resolution of the image is 20 20 10 m 3 Important tissue structures c an 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 rab bit 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 ver tical 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 mm 2 as shown in Figure 1 8 A Mechanical scan angles of 6have been achieved on both axes at more than 100 V. The y 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

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28 components pa ckaged 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 ) des ign 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 devic e 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 freq uencies 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 r ange 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 flowi ng 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 actua tors Outer torsion beam Gimbal Y X Control SMF FC lens Mirror AL package 5 mm

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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 curre nt the mirror scan angle c ould 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 wh en 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 mi rror 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, a s 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 di ameter and 12 mm length has been demonstrated with SD OCT. 3D images of finger tips were obtained at 18.5 frames/s ec 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 powe r 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 plat e size is 1.8 1.8 mm 2 and the device chip is as large as 10 10 0.2 mm 3 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 use d in a Fourier domain OCT and 3D human finger images were obtained.

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30 Figure 1 9 Electromagnetic ME MS OCT reported by Kim et al [71] A) SEMs of the ME MS mirror, B ) desig n 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 Epidermi s Dermis

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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. Pi ezoelectric 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 (P ZT). 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 piezoelectri c MEMS based OCT probe has also been reported [79] Piezoelectric MEMS mirrors with aperture sizes of 600 m 840 m and 840 m 1600 m have been fabricated. Mechanical scan angles up to 7 and resonant freq uency up to 1 kHz were measured for the mirrors. A prove of concept probe design is shown in Figure 1 11 B The 600 m 840 m mirror was used in a FD OCT system to demonstrate its imaging capability, and a 2D im age ( Figure 1 11 C) o f 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 angl e will complicate optical alignment and probe packaging. Furthermore, for piezoelectric

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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 Ele trothermal 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 coefficient s. 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 al so 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 the ir high fill factor. For the same mir ror 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 SiO 2 It was first demonstrated by Bhler et al in 1995 [81] Then Jain et al demonstrated 2D B C A

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33 micromirrors based on Al/SiO 2 bimorph beams and polysilicon as the heating material with rotation angle up to 40on a 1 1 mm 2 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 imagin g of porcine urinary bladders. 5 frames/s ec imaging speed was achieved with 2.9 mm 2.8 mm imaging area Yet d ue 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 3 7 [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 loo king endoscopic probe design, and E ) 2D OCT image of rabbit bladde r. A B D C E

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34 Figure 1 13 Electrothermal M EMS 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 rece ntly, 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 curle d 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 m diameter mirror apertur e 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) methodolog y 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 13 C. 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

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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 fo r endoscopic OCT scanners. Research Goals This research explore s 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 ea rly 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 disea se 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 sui table for both OCT imaging and photoacoustic imaging; photoacoustic imaging can achieve higher imaging depth at a

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36 cost of lower resolution, by combining the advantages of the two imaging modalities one imaging session would yield more information for diag nosis. 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 mi rrors 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 pola rization 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 prese nts 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 t he optical qualities of the pro bes 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.

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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 penetrati on 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 t o image the cardiovascular system with better contrast; while PS OCT targets regions with b irefringent properties. In this chapter the principle of low coherence interferometry will be introduced first. Then time domain and frequency domain OCT implementa tions 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 interfero meter. 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 divi ded 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 re combine 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 fiel ds of two beams of input light can be expressed as ( 2 1 )

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38 ( 2 2 ) In the equation s is the field amplitude vector; 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 expre ssed as ( 2 3 ) is the phase difference of the two signals. and are the individual intensities of the two input signals, and 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 The other requirements for obtaining a stable interference pattern are: first, the phase difference should not change with time, meaning ; and second, the initial phase difference 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 then the question could be simplified if we consider using a 50/50 beam splitter, the two beams will have the same amplitude and the same intensity As a result, the interference signal can be expressed a s

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39 is the optical path difference between the two arms, and the term is obtained here, because the light goes through the optical paths two times. Figure 2 1 Michelson i nterferometry. A ) Mi chelson Interferometer setup, B ) Interference pattern of wavelength The interference pattern of monochrome light source is shown in Figure 2 1 B It is a cosine wave with a period of 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, t he coherence length can be expressed in the fol lowing form : Only when the optical path B A Photodetector Beam splitter 0 x I

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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 t he 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 dom ain 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 deem ed 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

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41 Superluminescent Diodes (SLDs) are usually employed by time domain OCT, because they can offer high brightness, and low coherence length [96] Commercially available SLD s 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 m in air [97] Efforts have also been made to combine multiple SLDs t o 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 resol ution as high as about 3.7 m in air can be obtain ed [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 li nearly 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 s peed, 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 disadvantag es. 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 de lay 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

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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] L ater 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 r ealized 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. Th is 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 dispers ion in sample arm Figure 2 4 RSOD configurations. A) s ymmetric pulse shaping configuration B) f olded RSOD design employing double pass mirror [106] A B

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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 d epth scanning speed, Fou rier 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 informat ion 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 impl ementation, 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 impleme ntation 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, on e depth scan can be done in 250 s, 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

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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 : t omographic scan of the entire left circumflex coronary artery obtained with frequency domain OCT system (C7xR, LightLab, Watford, USA). A) longitudinal vi ew 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 ba sed 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 ca se 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 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 : A a B C D E

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45 ( 2 4 ) In this equation, is a constant, is the wave vectors of the scattered light, is the scattering factor. When assuming the scattering potential is constant over and direction, the equation can be simplified in the direction, a new constant could be used to include the integral over and directions, while is the distance from the sample surface to the photodetector ( 2 5 ) ( 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 interf erence power results in the autocorrelation of the scattering potential. ( 2 7 ) In an OCT system, when the reference mirror is placed in front of the sample with a distance and the sample thickness is T, shown in Figure 2 7 The strong reflection at the reference mirror can be expressed as where R stands for the reflectivity of the reference mirror. The scattering potential from an OCT system, is ( 2 8 )

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46 The inverse Fourier transform of the power of interference signal yields the auto correlation of the system scattering potential. ( 2 9 ) Figure 2 7 FD OCT signal detection schematic. is consist of four terms, the third term of the above equation is the tr ue reconstruction of the object structure, centered at In order for the 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 sam ple 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 numbe r 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

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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 fre quencies 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 t o 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 sweep ing period. Broadband source Beam splitter Reference mirror CCD array detector Latera l scanning PC DAQ Dispersive Grating A Sweeping source Beam splitter Reference mirror Photo detector Lateral scanning DAQ PC B

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48 SNR Advantage of FD OCT Without having to mechanically scan the referen ce 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 in which is the quantum efficiency, is the electron charge, and is the photon energy. This reference and sample arm signal power can be expressed by their power spectral density. ( 2 10 ) An assumption can be made here : the reference spectral density is the same as the light source spectral density ; while the sample spectral density is an attenuation of which can be noted as As a result, the signal power can be written as ( 2 11 ) There are three main noise sourc es in the OCT system, they are thermal noise, shot noise and relative intensity noise. The noise spectrum is expressed in the following equation. ( 2 12 ) In this equation, is the temperature in degrees Kelvin, is the value of the transimpedance amplifier feedback resistor, and is the

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49 coherence time of the source, which is 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 are the source spectral bandwidth and center frequency. BW is the signal passband, which is normally two times the carrier frequenc y spectrum. ( 2 13 ) When this signal is spatially separated onto two detectors, each detector obtain s half the original bandwidth, the signal power is the coherent summation of the signal over the spec trometer, while the noise power is the incoherent summation of the two. The new SNR is (2 14 ) As we can see here, SNR increased by a factor of 2. The above analysis can be extended to N det ectors, 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 : D oppler 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.

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50 Doppler OCT Doppler e ffect 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 sig nal collected by the photodetector is modulated by a carrier frequency According t o Doppler e ffect, if the imaged sample has a relative velocity to the incident light, the detected carrier frequency would change, and by detecting the changed carrier frequ ency, 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 ( ) to the incident beam with blood flow velocity of a relative velocity of could be generated, and blood vessels could be detected. Figure 2 9 Doppl er 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 Interference signal The spectrogram of a signal can be estimated by

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51 comp uting the squared magnitude o f the STFT of the signal as shown here : represents the i'th time window Each time window has a pixel size of 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 c alculating the spectrogram at carrier frequency 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 is obtained by calculating the centroid of the measured power spectrum. ( 2 15 ) This STFT method is effective. Yet its limitation lies in the fact that the minimum detectable Dopp ler 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 ch ange in consecutive A lines to compute the carrier frequency : Hilbert transform is used to obtain phase information. The detected interference signal is Hilbert transform can be applie d to obtain the complex signal as shown in equation (2 16) ( 2 16 )

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52 Figure 2 10 Consecutive A Line scans in Doppler OCT Doppler frequency shi ft is determined by the average of n number s 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. ( 2 17 ) ( 2 18 ) Polarization Sensitive OCT Certain tissues such as retina, meniscus, tendons etc. have birefringence property When such type s of tissue s are injured, this property will be damaged or lost. By detec ting 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

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53 as using Jones Matr ix, in which This linearly polarized light is then split into reference arm and sample arm by a 50/50 beam splitter, written as and 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, 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 a rm. 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 Assuming the birefringent axis of the sa mple 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 ( 2 19 ) Usin g Jones Matrix, the light coming back from the sample can be expressed as ( 2 20 )

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54 In this equation, is the sample's reflectivity, is the sample path length, which equals the distance from the beam splitter to the sample surface plus the depth into the sample ; as shown in this equation 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 photodetect ors separately. The horizontal and vertical signals at the photodetectors are shown in the following equations. ( 2 21 ) ( 2 22 ) The current detecte d at the photodectors are shown in equation ( 2 23 ) and equation (2 2 4) is the detector sensitivity, which is assumed to be the same for both

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55 detectors ; and is the path lengths difference between reference and samp le arms ( 2 23 ) ( 2 24 ) The cross correlation part of the signal is the interference signal, ( 2 25 ) ( 2 26 ) Assuming the light source has a Gaussian power distribution, is the spectral density, and is the FWHM, is the coherence length, integrat ing the signal over the light spectrum gives ( 2 27 ) ( 2 28 ) The signal at this point is demodulated at the c arrier frequency to eliminate the term,

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56 ( 2 29 ) ( 2 30 ) Let and by eliminat ing the constant term the signals can be expressed by the following equations. ( 2 31 ) ( 2 32 ) The reflectivity of the tissue can then be calculated from and the phase retardation is 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, and can be obtaine d, and can be calculated.

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57 Time Domain OCT S ystem 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 l ight 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 m axial resolution in air. The input light is divided into the reference arm and th e 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 fo r 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/s ec 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 22 Beam splitter Circulator RSOD Photo detector Endoscopic probe DAQ PC

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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 o f 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 ) Gre en light representing center wavelength B ) Blue light representing an additional wavelength. A B

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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 equ ation is is the grating constant, which equals to in our system, is the diffraction order, normally Take the green light as the center wavelengt h for example, as shown in Figure 2 14 A The RSOD is designed to give a diffraction angle, which means so we obtain In our system, so When the scanning mirror has a tilting angle of its optical path lengths change compare to tilting one way can be expressed by the following equation ( 2 33 ) Due to the double pass design and the direction in the graph, the actual scanning range is In our system, the scanning depth is designed to be 1.6 mm, so the tilting angle of the galvanometer should be Now consider dispersion, when there is a second wavelength as shown in Figure 2 14 B Its scanning range can be expressed as We can obtain based on the grating equation When is small, and its scanning range can be expressed as

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60 ( 2 34 ) The phase change induced is Knowing that can also be expressed as At this point, we can s ee 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 So by employing the grating based RSOD, dispe rsion can be eliminated, and a constant group delay linear to scanning angle is generated throughout the spectrum, which is When there is a misalignment between the center of the scanning mirror and its axis of rotation, a carrier fr equency 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 carrier frequency in RSOD can be exp ressed as This parameter is linear to misalignment 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 the misalignment When the system is perfectly aligned, that is there will be no carrier frequency, no matter what the scanning angle is. Therefore, the phase delay and group delay can be separate d and controlled individually. Spectrum of the carrier frequency induced by this method can be obtained

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61 by calculating the derivative of carrier frequency with respect to wavelength, as in 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 an d background were presented. Finally, the configuration and performance of the time domain OCT system employed in this study was introduced in detail.

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62 CHAPTER 3 ELECTROTHERMALLY ACTUATED MEMS MIRRORS A series of MEMS mirrors based on ele c trothermal act uation mechanism have been developed in BML aiming to be applied in end oscopic OCT imaging applications T his chapter first introduces the princi ple of electrothermal actuation and then goes on to introduce four generations of BML ele c trothermal MEMS mirr ors ; their design s 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 differe nt 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 he at 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 SiO 2 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 tilti ng angle can be calculated by as shown in Figure 3 1 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 SiO 2 which will cause the bimorph beam to

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63 bend downwards. The scanning angle is pro portional to the power, so it can be controlled through controlling driving voltage for the actuator. Figure 3 1 Eletrothermal bimorph structure. 1 st and 2 nd Generations Electrothermal MEMS Mir rors Several actuator or MEMS mirror designs based on this structure have been proposed i ncluding both 1D and 2D mirrors. The MEMS mirror shown in Figure 3 2 A has 1D scanning capability, it employed a series of b imorph 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 1 st generation mirror design Al Si SiO 2 Pt on actuation

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64 Figure 3 2 1 st generation MEMS mirror s A ) 1D MEMS mirror B ) 2D MEMS mirror. Despite the large scanning angle achieved b y the 1 st 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 artifact s in the imaging results. To solve this problem, the 2 nd generation mirror desi gn 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 p late 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 bi morph 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 vari ation 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

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65 Figure 3 3 2 nd generation: l ateral shift free (LSF) MEMS mirror design. A ) LSF bimorph actuator design B ) t op view model of 1D LSF MEMS mirror C ) SEM of 1D LSF mirror and D ) SEM of 2D LSF MEM S mirror. LSF LVD (3 rd ) and FDSB (4 th ) Electrothermal MEMS Mirrors The 3 rd and 4 th generations of electrothermal MEMS mirror designs were later proposed for endoscopic OCT imaging application s They are the Lateral Shift Free (LSF) Large Veridical Displace ment (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/SiO 2 bimorph beams and two rigid silicon frames. Pt is A B D C

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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 groun d which is connected to the middle pad. Electrical connection s 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 t he lateral shift and tilting during vertical displacement, t he 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, L 1 L 2 are the lengths of the two frames, and are the length s of the three bimorph beams. ( 3 1 ) Figure 3 4 LSF LVD eletrothermal actuator and mirror design. A ) s ide view of actuator and relationship between each element B ) t op view of the MEMS mirror. The mirror plate is held by four identical actuators. The size of the mirror pla te 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 m, leaving enough space for the mirror plate to either move downwards or tilt. This (a) Mirror Plate ACT 1 ACT 3 ACT 2 ACT 4 Bonding pads A B Mirror plate Bimorph III ( l 3 ) Frame II ( L 2 ) Bimorph II ( l 2 ) 0 0 Frame I ( L 1 ) Bimorph I ( l 1 )

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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 ap plied 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. Du ring 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 micro graphs 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 m 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 m 100 m A B C

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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 ) ver tical 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 pa ttern. E F C A B D

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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 use d for lateral scan. A ) SEM of the front side of the MEMS mirror, B ) SEM of the back side of the MEMS mirror s howing the Si ribs structure, C) l ayout of the MEMS mirror and D ) c haracterization 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

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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 suppor t 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 lat er used in the free space OCT setup for brain tissue refractive index measurement s 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 4 t h 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 bimo rph beams to have the same curling angle several geometrical requirements need to be satisfied, including a certain length ratio of the two beams 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 1 lists the optimized structure parameters that yield an initial displacement of 315

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71 Figure 3 8 FDSB actuator design. A ) s ingle S shaped bimorph B ) u se 2 S shaped to convert curling into vertical displacement and C) a ctual FSDB actuator structure. Figure 3 9 FDSB MEMS mirror. A ) s chematic 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 SiO 2 Al A B C

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72 Table 3 1: Optimized structure parameters of FDSB actuator Fixed Values Optimized Values Thickness (m) Adhesion SiO 2 : 0.05 Function Al : 1.0 Function SiO 2 (1st): 1.0 Heater Pt: 0.2 Insulation SiO 2 : 0.2 Function SiO 2 (2nd): 1.4 Adhesion Cr: 0.01 Section length (m) Total length L: 410 L 1 =252, L 2 =20, L 3 =138 Figure 3 10 Characterization of the FDSB mirror. A ) DC response of vertical displacement B ) DC response of optical angl e and C ) frequency response. A B C

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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. T he electric al connecting pads are formed on the surface of the device frame and t he mirror surface is on the other side W hen 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 i n Figure 3 9 The mirror aperture size is 1 mm 1 mm, and the chip size is 1.55 1.7 0.5 mm 3 The fill factor of this design is as high as 32%, and the device volume is only 1.3 mm 3 An initial downward displa cement of 340 m was achieved, and after a self annealing and burn in process, the displacement can increase to 375 m. The mirror plate is 25 m 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 m has been achieved, and by driving opposing actuators differentially optical scanning angle up to 23 has been obtained. The frequency response figure suggest s 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 introduce d 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 H owever, they have the disadvan tage of mirror center shifting during scanning. The second generation MEMS mirrors overc o me the center shift ing problem, and c an realize large piston motion movement. Yet, they could not realize tip tilt motion. The third

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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 employ ed 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 ge neration mirrors is the light we ight 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 r efractive index measurement s. 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. T his mirror design was employed in the fourth generation MEMS probe s

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75 CHAPTER 4 OCT ENDOSCOPIC IMAGI NG BASED ON MEMS This chapter focuses on endoscopic OCT imaging using probes built with MEMS mirrors. Four generations of probes will be introduced. T he optical and mechanical designs of the probes will be presented first; t hen the probe assembly and packaging schemes will be discussed ; after that, results of probe in vivo OCT imaging experiments will be d emonstrated ; i ssues 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 cou ld 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 4 th 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 pattern s 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 t ransversely to form 2D or 3D images. Figure 4 1 demonstrates 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 design s will be presented.

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76 Figure 4 1 Typical MEMS probe design. Probe Optical D esign The three major optical components in an OCT probe are single mode optical fi ber, 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 position al 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 lig ht beam from fiber

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77 Single mode f iber (corning SMF 28) is chosen to avoid modal dispersion in light transmission. Core diameter of the fiber is 8.2 m, its effective refractive index is 1.4677 at 1310 nm, its numerical aperture is 0.14, its cladding diameter is 1250.7 m, and the coating diameter is 900 m. 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, th e 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 lens es is easier to package in smal ler spaces. The GRIN lens has a center axis refractive index of 1.616 4, 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 m, 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 diameter s and working distance s 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.

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78 The first 3 generations of endo scopic OCT probes were designed and assembled based on the 2 2 0.5 mm 3 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 1 st generation probe a two piece design using Lucite as material was chosen. One piece i s used to hold the MEMS mirror, while the other piece i s for the other optical components, including fiber and GRIN lens. The MEMS piece has a protruding part, and the other piece h as 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 su ccessful ass emblies. 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 piece s were

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79 registered together through the slot. After all the above steps were completed, the probe was inserted i nto flexible, biocompatible, transparent fluorinated ethylene propylene (FEP) tub ing with 5. 8 mm outer diameter A finished probe is shown in Figure 4 4 C. Figure 4 4 1 st generation probe design. A) fabricated probe body in Lucite B ) e lectrical 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 d esign 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 p robe. 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 2 nd generation design us ing 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 negli gible. 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

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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. Th e diameter of the prob e mount was 5 mm; after being inserted in to the FEP tub ing the packaged diameter wa s 5.8mm. Figure 4 5 2 nd generation probe design. A ) 3D model of the assembled probe B ) assembled probe in stain less 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 a re 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

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81 connection the MEMS mirror to the PCB. This design belongs to the 3 rd generation, it has successfully reduced the diameter of the probe to 2. 8 mm. 3D views of the probe are demonstrated in Figure 4 6 Figure 4 6 3 rd generation probe de sign. 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 3 rd generation design, a new 4 th generation prototype based on the DFSB mir ror 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 e lectrical connection has been made easy due to the special design of DFSB mirrors which have their connection pads on the backside of the m irror frame. In A B C D

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82 this design, n o wire bonding process is necessary therefore, a lot of space could be saved and f urther 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 3 rd generation design, this probe uses flexible PCB for electrical connection. The difference is t hat 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 flexib le 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

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83 probe is then packaged with a transparent FEP tub ing which has an inner diameter of 2.5 mm and an outer diameter of 2.8 mm. Figure 4 8 DFSB probe design. A) f lexible PCB design, B) zoomed in view of the PCB head with bonding pads pattern, and C ) packaged probe. In Vivo Experiments The 2 nd and 4 th 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) intraperiton eally. 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 2 nd generation probe, the actuator pair on the circumferential direction was used for fast axi s scan, driven by 0~4 V differential ramp voltages at 1.25 Hz; and the actuator pair on the longitudinal direction wa s 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

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84 in 40 s for one 3 D reconstruction. This generates a transverse area of ~2.3 2.3 mm 2 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 th e 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 m. 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 pu t 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

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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 4 th generation probe, mouse ear and tumor were imaged. For this imaging session, h uman breast cancer cells, 4T1 mammary carcinoma cell line, were injected into the subcutaneous tissue of the f emale athymic (nu/nu) nude mouse with a total volume of 10 And t he 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 f or the 2 nd 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. T he tumor edge can be clearly seen in 3D images and 2D images demonstrated obvious diff erence between normal and tumor tissue areas, due to strong penetration loss inside the tumor cell. E D E D cc cc C 1 m m A B

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86 Figure 4 11 In vivo imaging results. A ) mouse under ear imaging session B ) 2D image of mouse ea r, C) 3D image of mouse ear, D) probe position of tumor imaging, E) inj ected tumor under microscope, F G ) 3D mouse tumor imaging results, and H J ) 2D images of tumor margin, locati ons 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 te lecentric scan, circled in Figure 4 1 2 To understand these issues, Code V (Optical Research Associates) simulation s w ere conducted ( e ) 1 mm Probe tub ing 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

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87 Figure 4 1 2 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 pr obe. 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 defined GRIN material with center axis refractive index of 1.616 4, 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 diame ter 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, t he b est 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

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88 calculated using Gaussian beam tracking. Surface properties and thickn esses 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 s imulation results, a setup was built using 3D movement stages to simulat e 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 g lass slide with a 1951 USAF resolution pattern. The r esolution of the system wa s determined by the finest pattern distinguishable from the OCT images. And the focal length wa s recorded as the center of the clearest image range. GRIN lens Fiber output FEP tub ing MEMS mirror 50 m B A

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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 w ill 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 simulation s and experiments were condu cted, and their results showed good agreement with each oth er. 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 change s more than 3 mm. The d esired working distance of about 5 mm c an be obtained when L1 is controlled at approximately 0.034 mm; and as long as L1 is smaller than 0.08 mm, the probe w ill have an acceptable focal length of larger than 3.5 mm meaning the beam focus will be outside t he 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 m

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90 agreement w ith simulation results when spot size is at 60% of the maximum intensity from the simulation. The s pot size is larger than 30 m within the acceptable working distances for this probe design. Figure 4 15 Effects of the distance between the fiber and the GRIN lens. A) s chematic of the testing setup B) effect on focal length, and C) effect on spot size. C B A

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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 curv ature (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 T he 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 (perfec tly flat mirror plate) radius of curvat ure 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

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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 smal ler ; this may seem desirable for the sake of higher resolution; however, t he drawback is that astig matism becomes more and more serious as the radius of curvature decreases. When there is astigmatism in the system, l ight in the sagittal direction and the meridional direction will be focus ed at two different planes, the focus for this type of system is t hen 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 4 th generation FDSB MEMS mirror and the mirror radius of curvature is about 10 m. Figure 4 18 A is the comparison between the tested fo cal 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 c ould distinguish slightly better than FWHM of the spot diameter with the MEMS scan system.

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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) d istance between fiber and GRIN lens vs. focal length, B) d istance between fiber and GRIN lens vs. spot size (simulation result is FWHM spot size at 10m radius of curvature, experimental result is best observed r esolution pattern size) Effect of FEP T ubing 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 effect s of this tubing on the o ptical 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

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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. F rom 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 mat erial, 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 differe nt 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 m at most. Then, the MEMS mirror wa s 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 increas ed 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.

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95 Figure 4 19 Effects of FEP tubing on optical beam from the GRIN lens A) s chematic 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

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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

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97 when the mirror rotates in plane, the angle change of the reflection beam will be twice of the rotation al 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. A ccording 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 curvat ure 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 cente r position was calculated and then entire A line was shifted accordingly to match up with the center position. Specifically, i n the OCT system, controlling signal for the MEMS mirror is outputted through a data acquisition card (USB 1408FS, Measurement Com puting). 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 me asured at about 2 mm from the MEMS center the entire imaging depth was 1.6 mm Therefore, for the K th 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. ( 4 1 )

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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 in the radial direction and the beam scan angle could be calculated by equation 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

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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 de sign 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 probe s. 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 m. The working distance and resolution were not significantly affected by the 4 th generation MEMS mirror, due to its large radius of curvature. FEP tubing introduces slight astigmatism, when MEMS mirror is no t perfectly flat. And the resolution will not change much at different scan angles during imaging. A B

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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 p hysical thi ckness and optical thickness of the brain tissue slices. The large aperture LSF LVD MEMS mirror was employed for transverse scan in order t o 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 ti ssue. 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 measur ing 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 d iagnosis can be made based on the refractive index difference between normal and malignant tissues [142] [143]

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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. Oth er 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 kid ney, 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 use d OCT image s of a flat substrate and tissue boundaries and extracted the RI by compar ing the optical and physical thickness; and the second method utilize d focus tracking, in which the movements of the reference and sample arms were recorded and the RI was calculated based on the distance s they move d 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 g roup 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 tissu es due to the chromatic dispersion

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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. Meas urement 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 va riation 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 c ell 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 elastogra phy, 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 t he sample for lateral scans, the light weight large aperture LSF LVD MEMS mirror introduced in Chapter 3 was used ( Figure 3 7 ). The o utput light from the single mode fiber (SMF 28) was collimated with a fiber collimator (Thorlabs, F260FC C); and the collimated beam size wa s 2.8 mm in diameter. The light wa s then focused using a lens with diameter of

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103 12.7 mm and focal length of 39 mm. A MEMS mirror wa s placed approximately 26 mm behind the focusing lens, with 45 to the incoming beam. The beam diameter at the mirror plate wa s about 0.94 mm, and the 45 cross section wa s 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 wa s ~23 m. 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 st age movement, even small relative movement of the tissue would cause image artifacts, and by employing a MEMS mirror, mo tion 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

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104 compression. Rat surgery was conducted in accordance with the NIH guidelines on the use of anima ls 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 eut hanasia, 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 m 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 O 2 saturated artificial ce rebrospinal fluid (aCSF, between approximately 35 37 C. The a CSF was continuously bubbled with 95% O 2 and 5% CO 2 gases and 0.5 mM L Glutamine (Invitrogen Co., CA) and 1% penici llin 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 condition s ; cell viability was determined by measuring neuronal death or degener ation 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 i s calculated by dividing optical thickness b y its physical thickness i.e ( 5 1 )

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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% O 2 and 5% CO 2 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 ge ntly 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 m was measured to be between 496 to 498 m 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. Media l sections from excised rat brains were sliced into coronal sections of 300 m thickness.

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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 dista nce 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 ph ysical 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 minu tes for each ser ies 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 dur ing the experiment, B ) OCT im age of the brain tissue slice, and C ) OCT image of air after brain tissue is removed. A B C tissue thickness air t hickness

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107 Figure 5 4 Measurement of optical and physical thickness of various anatomical r egions in rat brain tissue slices. A1 E1) meansurement of t he optical thicknesses (t optical ) of each target region and A 2 E2) measurement of the physical thicknesses (t physical ). A 1 A 2 B1 B2 C1 C2 D1 D2 E1 E2

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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). A 1, B 1) no compression, A 2 B2 ) 20%, A 3, B 3) 40%, A 4, B 4) 60% and A 5, 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 delin eated. 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 w as 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 regi on ranged from 0.007 to 0.016. The RI of the aCSF was 1.342 0.007. This w as ~2.7% lower than the average RI over all the regions measured in the brain tissue slices, which was 1.380. A1 200 m A2 A3 A4 B1 A5 B2 B3 B4 B5

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109 Figure 5 6 Measured RI in various anatomical regions in brain and aCSF. 10 samples were measured at each region. Bars correspond to 1S E (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 defo rmation 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% w as measured for the grey matter, while an average 90% increase was measured for the white matter.

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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, R I 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 tha n 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 1.4070.015. No statistical differences in RIs of grey matter regions (cerebral cortex and putamen) were f ound, and the average RI of all grey matter regions was 1.3690.014. 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 dif ferences 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,

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111 because of the large water con tent 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 m wavelength [149] 1.371 over a range of spatial frequency of 0.06 to 0.6 m 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 ti ssue 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 bot h 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 increas ed 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

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112 constituents comp ared 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 wh en 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 m (1SE=14 m) and 254 m (1SE=22 m), 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 r esulting in thickness changes. Ho wever, 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 t hat 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 tr ansport 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

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113 not be significantly compromised during this short period of time. Cells may also be damaged due to lo ading 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 no n 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 i ndentation 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 R I 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 testin g of tissues under compression or indentation tests may also benefit from this study.

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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 t his 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 adde d to the existing OCT configuration While the major difference for implementing Doppler OCT from the basic system is signal processing algorithm. In this study, e x vivo PS OCT animal imaging using MEMS probe has been realized; while Doppler OCT imaging re sults are obtained by using translational stage for lateral scan s 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 its Stokes vectors are defined a s

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115 ( 6 1 ) ( 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: polarizat ion controller; BS: 2 x 2 beam splitter; RSOD: rapid scanning optical delay line; PBS: polarization beam splitter; D A and D B : balanced photodetectors. Infrared light coming from the light source goes through a polarizer first in order to obtain linearly p olarized 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 li ght goes through the system. When a voltage is applied to the modulator, a phase shift

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116 can be induced, and polarization state will change from linear to circular. The interference signals from these two polarization states are used to calcula te 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 polarizatio n is used. The input light is the reference signal reaches PBS1 can be written as and the reference signal at PBS2 can be written as ; similarly the sample signal at PBS1 is and the sample signal at PBS2 is The reference and sample signal s 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 D A and D B the DC signal is cancelled at this point, and the detected signals are shown in equation (6 5). ( 6 3 ) ( 6 4 ) ( 6 5 )

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117 Signal processing using Hilbert transform gives the complex signal as shown below. ( 6 6 ) With th ese results, a set of stokes parameters of linearly polariz ation can be obtained using the following equation. ( 6 7 ) During the next time period, voltage is applied on the polarization modulator to induce circula r polarization. A prime is used to distinguish circular polarization from linear polarization The input circular ly polarized light can be represented by Similar to the linear case, the reference light reaches PBS1 is and the reference light at PBS2 is The sample signal s at PBS1 and PBS2 are and Again the same polarization orientation will interfere with each other, and the interference sign als are ( 6 8 )

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118 ( 6 9 ) The detected signals from circular ly polarized input light are ( 6 10 ) The complex signal after Hilbert transform can be express as ( 6 11 ) The signal Stokes vectors with circular polarization input light can be calculated by the following equation ( 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 represen ting 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

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119 cut open ( Figure 6 2 ) to expose the menisci. The 2 nd generation OCT probe was employed for thi s 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

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120 A lateral scanning area of 2.3 2.3 mm 2 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 im ages, 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 r econstructed 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 tub ing Doppler OCT System Doppler OCT Principle To realize Doppler OCT, no hardware modification is required for the original time domain OCT system. Only an addition al 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 c ould be calculated from the equation : Hil bert transform is used to obtain the phase information once signal is collected by the data acquisition card. During signal processing Hilbert transform is realized by fast Fourier transform (FFT). It could be done ac cording to the following procedure: In this method, Hilbert transform is realized in the frequency domain F irst a Fourier transform is conducted and then the signal is multiplied with the function

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121 A b and 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 expe riment s on imaging mouse blood vessels w ere conducted in our lab. T ransverse scan s w ere 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 result s of Dopp ler 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 v ivo Doppler imaging. A) s ample mouse during imaging session, B) i ntensity image through the window chamber and C) Doppler image through the window chamber. Blood Vessels A C B

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122 Summary In this chapter, detailed implemen tations of PS OCT and Doppler OCT at the BML were introduced. PS OCT was successfully demonstrated with the 2 nd generation MEMS probe. 2D and 3D OCT images of canine meniscus were presented. Improved contrast of Doppler OCT for rat blood vessels was also s uccessfully demonstrated.

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123 CHAPTER 7 MEMS OCT/PHOTOACOUSTIC IMAGING So far we have discussed For highly scattering tissues, the penetration depth of OCT could be limited. It is therefore desirable to u se 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. I t can normally reach 3 5 mm de e p in tissue, with about 100 m 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, f irst the background and principle of PA imaging will be introduced. Then the photoacoustic imaging s ystem 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 Backgroun d PA imaging is based on the PA effect, which describes the generation of acoustic wave s by the inciden ce of optical energy. In the case of PA imaging, when laser output is incident on to the tissue, the absorption of the optical energy cause s the tissue t o have thermal expansion, and subsequently, emitting broadband ultrasonic waves. T he time of flight, amplitude etc. of the ultrasonic waves provide information regarding

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124 location, strength and dimension of the acoustic source, therefore tissue reconstructi on 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 a ttention recently. Researchers have demonstrated PA imaging applications in breast cancer, brain vasculature, skin disorder and arthritis, etc. [164] [171] For endoscopic photoacoustic imaging, met hods 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 li mits 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 3 rd generation LSF LVD MEMS mirror is designed an d 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 a ttached to the edge of the probe; it is a one element ultrasound detector, which is fabricated with a 110 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. W hen light comes in, it first incident s 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, t he fo ur actuators are driven by 0 ~ 4 V differential ramp voltages and optical scan angles of

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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 mm 2 could be obtained. The scanning beam size i s 1 1 mm 2 determined by the MEMS mirror aperture size. Figure 7 1 Endoscopic photoacoustic probe. A) schematic of probe, B ) 3D model and C ) a ssembled 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 (N L 303HT from EKSPLA, Lithuania) ; it generate d pulses of 20 ns duration at 10 Hz. Two functi on 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. S ampling 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 A C

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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/cm 2 which is below the American National Standards Institute safety limit of 20 mJ/cm 2 For 2D images, 50 p oints of signal w ere obtained in each direction; the scanning frequenc y for fast scan was 100 mHz, and for slow scan, it was 2 mHz. T he scanning time was limited to 250 s d ue to the 10 Hz repetition rate of the laser. During the experiments, ultrasound gel was applied bet ween 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

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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 reconstructe d 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) pho tograph o f the imaging target B) D) coronal, sagittal and cross section views of the recovered 3D imag e and E) 3D rendering of the recovered image. Figure 7 4 In vivo photoacoustic experiment. A) imaging tar get B) reconstructed blood vessel image. MEMS Probe for B oth 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 m ode 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 connect ed to the OCT system and the PA A B C D E A B reconstructed blood vessel

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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 4 th 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 4 th 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) s chematic o f 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 V ivo Experiment The assembled probe was employed in ex vivo experiments to de monstrate its capability for both OCT and PA imaging. The same procedures as that described in A B MEM S mirror transducer GRIN lens GRIN lens transducer

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129 Chapter 4 for the 4 th generation probe was adopted for this experiment as the same type of MEMS mirror was employed. From Figure 7 6 l ayered 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 t hen a side view endoscopic probe that has the ability for both OCT and PA imaging was introduced. 2D and 3D OCT imag es of a m ouse ear were obtained with the prototype MEMS probe. A E D D E cc cc C 200 m B

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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, f our generations of ele c trothermal ly 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 4 th 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 c learly separated from healthy tissue structures. Secondly, MEMS mirror based free space OCT for refractive index (RI) measurement in brain tissue slices has been demonstrated. RI s 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 disease s Information on RI change under compression can also help the mechanical p roperty tests of soft tissues based on elastography.

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131 Thirdly, two types of functional OCT were demonstrated. They are polarization sensitive OCT and Doppler OCT. The implementation s of these two types of OCT systems help improved the image contrast. PS OC T 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 modalitie s, 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 W ork T he 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 achievement s of this work for MEMS OCT probes to be widely used clinically, more work needs to done. Firstly, m ore robust assembly procedure s 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 f lat tubing and adding astigmatism correction lenses are two potential resolutions that could be explored.

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132 Thirdly, an imaging rendering algorithm that takes into account the non telecentric scan effect and non linearity of the MEMS scan should be develope d 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 s ide view probe was used to image canine meniscus ex vivo in this study A front view probe is more suitab le for imaging the meniscus in live animal s 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 b rain tissue slices on RI and RI changes in response to compression should be studied in the future. Lastly, t he first prototype probe for OCT and PA imaging has a relatively large diameter Fo r clinical use, new designs with smaller sizes are desirable.

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133 APPENDIX PUBLICATIONS The following is a list of archival journal and c onference 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 b ased e ndoscopic o ptical c oherence t omography", Int. J. Opt 2011, 825629, ( 2011 ) 3. J. Sun, S. Guo, L. Wu, L. Liu, S. Choe, B. Sorg, an d H. Xie, "3D i n v ivo 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 b ased 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 v ias", J. Biomed. Opt. 16 026006 (Feb 01, 2011). Conference papers and abstracts: 1. to be presented at Biomedical optics (B IOMED 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 v itro r efractive i ndex m easurement of a cute r at b rains u sing o ptical c oherence t omography ", presented at the 2010 Biomedica l 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 (2 010)

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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 p roperty c hanges of a cute r at b rain t issue s lices a s a f unction o f c ell v iability", 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 o ptically b ased i ndentation t echnique for t hin s oft b iomaterials ", 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 d imensional o ptical c oherence t omography b ased on a h igh f ill f actor m icroelectromechanical m irror 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 im aging probe based on a high fill factor MEMS mirror and wire bonding Micro Electro Mechanical Systems (MEMS), 2011 IEEE 24th International Conference on ( Cancun, Mexico, 23 27 Jan. 2011), pp.33 3 6. 10. D. Wang, L. Fu, J. Sun, H. Jia, and H. Xie, "Design o ptimization and i mplementation of a m iniature o ptical c oherence t omography p robe b ased on a MEMS m irror", Proc. SPIE 8191 81910M (2011) 11. W. Liao W. Liu L. Xi J. Sun Y. Zhu L. Wu H. Jiang and H. Xie Miniature p hotoacoustic i maging p robe u sing MEMS s canning m icromirror ", Optical MEMS and Nanophotonics (OMN), 2011 International Conferenc e on (Istanbul, Turkey, 8 11 Aug. 2011 ), pp.115 116

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150 BIOGRAPHICAL SKETCH Jingjing Sun was born in Tianjin, China in Janua ry of 1985. She enrolled in Tianjin University in Tianjin, China in the fall of 2003. F our years later, in July 2007 she received her B degree in information engineering After graduation, she joined the Biophotonics and Microsystems Lab at the U niversity of Florida and started working with Dr. Huikai Xie on MEMS based OCT imaging research I n December 2010, s he obtained her Master of Science degree in electrical engineering And she is expected to graduate with her Ph.D. degree in electrical e ngineer ing in May 2012.