High Resolution Optical Coherence Microscopy
by
Aaron Dominic Aguirre
B.S.E., Electrical Engineering
University of Michigan, Ann Arbor, 2000
Submitted to the
DEPARTMENT OF ELECTRICAL ENGINEERING AND COMPUTER SCIENCE
in partial fulfillment of the requirements for the degree of
MASTER OF SCIENCE
at the
MASSACHUSETTS INSTITUTE OF TECHNOLOGY
February 2003
0 Massachusetts Institute of Technology 2003
All rights reserved
Signature of Author
Department of Electr aI Engineering and Computer Science
February 2003
Certified by
Professor James G. Fujimoto
Thesis Supervisor
Accepted by
Chairman, Departient committee on Graduate Students
MASSA CHUSETTS INSTITUTE
OF TECHNOLOGY
MAY 12 2003
High-Resolution Optical Coherence Microscopy
by
Aaron Dominic Aguirre
ABSTRACT
Optical coherence microscopy (OCM) is a technique that combines the high transverse
resolution of confocal microscopy with the coherence gated, heterodyne detection of optical
coherence tomography. By combining confocal spatial rejection and coherence gating to remove
unwanted scattered light from images, OCM can yield improved contrast and greater imaging
depths than standard confocal microscopy. Real-time, in vivo OCM has been demonstrated for
cellular imaging.
To take full advantage of the improved axial sectioning provided by coherence gating, OCM
systems must be designed to support large optical bandwidths available with femtosecond laser
sources. Construction of real-time, broadband OCM imaging systems has previously been
limited by the availability of high-speed, broadband phase modulators. Earlier work has used
either a fiber-stretching piezoelectric modulator, which limits speed, or a waveguide electrooptic phase modulator, which limits the optical bandwidth of the system. Furthermore,
waveguide devices are commercially available only at select wavelengths. This thesis discusses
the demonstration of a novel, broadband OCM system that enables real time imaging of cellular
structure in highly scattering tissue. The system integrates a high-resolution OCT system with a
reflective grating phase delay modulator and a fast scanning confocal microscope. Grating phase
delay scanners have been developed and demonstrated previously for high-speed OCT imaging
and for phase modulation. The novel reflective geometry demonstrated here enables OCM
imaging with large bandwidth, providing coherence gates of only a few micrometers. Moreover,
the flexible OCM system design can readily be implemented at wavelengths that were previously
inaccessible for OCM imaging.
The broadband system is used to demonstrate a new operating regime for real-time, in vivo
OCM imaging of cellular structure in human tissues. Combined coherence and confocal gating
is shown to relax microscope design constraints imposed by confocal microscopy. In particular,
a short coherence gate is used to enhance weak confocal sectioning, thereby enabling cellular
imaging in situations when confocal microscopy alone would be inadequate. The results
demonstrated offer promise for cellular imaging in clinical applications that require probe
technology unsuitable for confocal imaging. As a first step toward such clinical applications of
OCM, a compact handheld imaging probe is developed and demonstrated.
Thesis Supervisor:
James G. Fujimoto
Professor of Electrical Engineering and Computer Science
3
Acknowledgements
I would like to thank my thesis advisor, Professor James Fujimoto, for providing the
guidance and the resources necessary to complete this work. His keen scientific insight, tireless
attention to detail, and careful mentoring inspire me personally and professionally and I am
grateful for the opportunity to work with him. I would also like to thank my colleagues in the
Ultrafast Optics Group at MIT. Together they create an exciting and supportive environment for
scientific research.
This work could not have been completed without the technical support and friendship of
several people. Pei-Lin Hsiung helped me to get started in the lab and worked closely with me
on much of the system development. Her technical expertise made the work much easier and her
optimism and sense of humor made it much more enjoyable. Tony Ko provided continued
technical advice and support with software and system details and has become a trusted friend.
His work ethic and selflessness are admirable. Stephane Bourquin provided much technical
advice and will remain a close friend after our work together ends. Ingmar Hartl designed the
electronic receiver and was a pleasure to work with during my first year in the group. Drew
Kowalevicz and Rohit Prasankumar unselfishly provided help with laser alignment and have
become good friends as well. The technical contributions and support of Paul Herz, Aurea
Tucay, and Alphan Sennaroglu are also very much appreciated.
I gratefully thank my friends from MIT and Harvard for providing the necessary diversions
to keep me sane during the past two years. In particular, I acknowledge Joaquin Blaya, Jenny
Mu, Kevin King, Roxanna Webber, and Todd Coleman for their support. I also thank all of my
friends from Michigan, especially Axel Berny, Chethan Gangireddy, Gar Dewey, Nita Parekh,
Sarah Dehaan, Seth Myers, and Vaishalee Padgaonkar.
I endlessly thank my parents for all of their love and support. They have sacrificed so much
for my brothers and me and I will always define myself by the things I learned from them.
Finally, I thank my brothers Andy and Derek. They are my best friends and deserve special
recognition for putting up with me while I wrote this thesis.
To those who know...
STBDFTBH
4
Contents
Abstract.......................................................................................................................................3
A cknow ledgem ents ....................................................................................................................
4
Table of C ontents .......................................................................................................................
5
Chapter 1: Introduction
9
1.1
M otivation ......................................................................................................................
9
1.2
H igh R esolution Im aging in Tissue ........................................................................
9
1.2.1
1.2.2
1.2.3
1.2.4
U ltrasound ......................................................................................................
Optical Coherence Tom ography ......................................................................
Confocal M icroscopy ......................................................................................
Tw o-Photon M icroscopy ..................................................................................
10
11
13
15
1.3
O ptical C oherence M icroscopy ...............................................................................
16
1.4
Previous W ork on O CM ...........................................................................................
19
1.5
Scope of Thesis ..............................................................................................................
21
R eferences ..................................................................................................................................
23
Chapter 2: Optical Coherence Microscopy
29
2.1
O verview ........................................................................................................................
29
2.2
Scattering in Biological Tissues ...............................................................................
30
2.3
C onfocal M icroscopy ...............................................................................................
32
2.3.1
Im age Form ation in Confocal M icroscopes ...................................................
32
2.3.2
Lateral Response .............................................................................................
37
2.3.3
2.3.4
2.3.5
A xial Response and Sectioning .....................................................................
Effect of Aberrations ......................................................................................
Effect of Finite Detector Size ........................................
38
40
40
5
2.3.6
2.3.7
2.4
2.5
44
2.4.1
2.4.2
2.4.3
2.4.4
2.4.5
2.4.6
2.4.7
44
46
48
49
50
51
52
Interferom eter Analysis ...................................................................................
Coherence Gating .............................................................................................
Effect of Group Velocity Dispersion .............................................................
Detection Electronics ...................
.........................................
Noise Sources.............
................
.........
................
System Sensitivity . ....
....
.........
......... ......
Dual Balanced Detection .................
......
...
.............
Combined Confocal and Coherence Gating
.................
...........
R eferences
63
................................... 64
Grating Conventions and Notation
Phase and Group Delay Equations ...............................
65
Dispersion Compensation
....... .......................................................
69
.............................
........................ .....
3.1
Overview ....................................
3.2
Requirements for In Vivo Cellular Imaging
3.3
Broadband Light Sources
..................... 73
...
............................
73
.......................... ................... 74
Semiconductor Superluminescent Diode Laser Source at 1300 nm. .......... 74
Modelocked Ti:A120 3 Femtosecond Laser Source at 800 nm .................. 75
Fiber Broadened Femtosecond Laser Sources at 1064 nm and 1250 nm ..... 77
Interferometer ............................................
3.4.1
70...............
70
73
Chapter 3: OCM System Development and Characterization
3.3.1
3.3.2
3.3.3
53
Heterodyne Signal for Combined Gating......................... o ...............
53
54
Depth of Field and Transverse Resolution ...............................................
Optical Coherence Microscopy for High Resolution Imaging ......
........ 55
Path Length Scaling with Focal Position in Tissue ....................
57
Enhanced Gating Effects in Scattering Media ......
............
...... 58
Operating Regimes for OCM ...........
.
....................................
60
Phase Delay Line Modulator ............................................
2.6.1
2.6.2
2.6.3
3.4
41
41
Low Coherence Interferometry ..............................................................................
2.5.1
2.5.2
2.5.3
2.5.4
2.5.5
2.5.6
2.6
Fiber Optic Confocal M icroscopes ................................................................
Scanning Confocal Microscope Designs .......................................................
Spectral Transmission Measurements ....................
6
..................
.
78
............. 78
Polarization Control ........................................................................................
80
Reflective Grating Phase Modulator ......................................................................
80
M odulator D esign ..........................................................................................
M odulator Characterization ............................................................................
81
83
Sample A rm O ptics .......................................................................................................
85
3.4.2
3.5
3.5.1
3.5.2
3.6
3.6.1
3.6.2
3.6.3
3.6.4
3.6.5
3.6.6
M icroscope Objectives ....................................................................................
Reflective Microscope Design for Finite Tube Length Objective ..................
Close-Coupled Scan Design for Infinity Corrected Objective ........................
..........................
Performance Under Broadband Illumination ....
Combined Gating Effects ..........................................
Compact Handheld Imaging Probe .................................................................
3.7
Receiver Specifications ......................................
3.8
Image Acquisition and Processing
3.8.1
3.8.2
3.8.3
3.8.4
....
...........................................
95
96
...................................
Timing and Synchronization .......................................
Softw are Interface ..........................................................................................
Zipper Effects ............................ ... ...............................................................
Sam pling Criterion ...........................................................................................
3.9
System Sensitivity M easurement ...............................................................................
3.10
Axial Resolution Measurement.............................................102
......................
References ...............................................
85
86
88
92
93
94
97
98
99
99
100
................ 105
107
Chapter 4: In Vivo Imaging Results
.... ................. .............
.............
.........
.........................
107
4.1
O verview ....................
4.2
Imaging of an Animal Model: Xenopus laevis Tadpole ........................
107
4.3
In Vivo Imaging of Human Skin .............................................................................
110
Exposure Limits for Microscopy ....................................
Tissue Stabilization ..............................................
Imaging with Short Coherence Gate ..................................
Im aging Depth M easurem ent ..............................................................................
111
112
112
114
4.3.1
4.3.2
4.3.3
4.3.4
4.4
Preliminary Imaging Results with a Handheld Probe at 1300 nm .......................... 116
7
References ..................................................................................................................................
Chapter 5: Summary and Future Work
8
118
121
Chapter 1
Introduction
1.1
Motivation
Advances in genetics and cell and molecular biology have both enabled and necessitated an
understanding of human disease at the cellular and molecular levels. This has driven a
corresponding desire to develop high-resolution medical imaging techniques that provide
diagnostically useful information about the microscopic state of tissues. Excisional biopsy and
subsequent histologic examination is the current standard for assessment and definitive diagnosis
of disease at the cellular level [1]. Biopsy and histology, however, are invasive and sometimes
high risk and therefore not conducive for widespread screening for early stage pathologies.
Furthermore, histologic sectioning is time consuming and cannot provide real time analysis of
tissue state. A need exists for minimally invasive, high-resolution imaging techniques that can
provide real time information about tissue microstructure.
A technique known as optical coherence microscopy (OCM) has potential to help address
this problem. This thesis will discuss development of enabling technology for OCM and will
assess the potential of this modality for in vivo imaging of human tissues.
1.2
High Resolution Imaging in Tissue
Visualization of tissue microstructure requires imaging resolution that corresponds to the size
scale of the structures themselves. Such whole body clinical imaging techniques as x-ray
computed tomography (CT) and magnetic resonance imaging (MRI) provide essential tools in
assessing features as small as 500 um to 1 mm, but this resolution is not sufficient to image
important cellular structure in tissue. The entire thickness of epithelial cell layers which coat the
body's internal and external surfaces and cavities is not usually greater than a few hundred
microns and can be as thin as a single cell layer. Individual healthy epithelial cells can vary in
diameter from about 2 - 25 um with their nuclei typically only 50% or less of this dimension.
Capillary blood vessels range in diameter from 3 um to about 30-40 um while small lymphatic
vessels and glandular structures are not typically larger than a few hundred microns in dimension
[2].
Research on high resolution CT and MRI techniques have improved resolution to a
desirable scale in the lab, but these techniques are physically impractical for human tissue
imaging. Research MRI machines capable of resolving structures smaller than 50 um require 414 Tesla magnets with prohibitively small bore sizes and are limited to long acquisition times to
improve intrinsically poor signal-to-noise ratio [3-5]. Soft x-ray microscopy research has
reached resolutions down to tens of nanometers, but this technique also suffers from restricted
field of view and long image integration times [6]. Furthermore, x-ray techniques have the
additional disadvantage of using ionizing radiation, making prolonged imaging of living
biological specimens difficult [7]. Hence, development of modalities for fast imaging of tissue
microstructural features has shifted to acoustic and optical techniques where resolution on the
9
order of the wavelength of emitted waves can be achieved with non-ionizing radiation. Progress
in ultrasound imaging and in optical imaging has generated hope for a clinically useful, highresolution technique.
1.2.1
Ultrasound
Ultrasound imaging is a direct, non-reconstructive form of imaging. A high-frequency
acoustic wave is launched into tissue from a piezoelectric transducer and echo time delays of
reflections from tissue scatterers are measured electronically. Lateral spatial resolution is
determined by transducer focusing characteristics and is approximately equal to the acoustic
beam width at the location of the scattering object. Beam width is determined by diffraction
from the transducer geometry and scales with wavelength of the radiation. Hence, the lateral
resolution improves with higher frequency sound. Axial gating in the direction of wave
propagation is achieved by pulsing the source wave, and axial resolution is thereby determined
by pulse width. Higher frequency sources allow for improved axial resolution as well, since
shorter pulse durations are possible. A fundamental tradeoff exists between resolution and
penetration depth, however, because tissue absorption increases with increased acoustic
frequency [8]. Contrast in ultrasound images is a function of mechanical properties of tissue and
the technique therefore offers a complementary set of information to optical imaging modalities.
Ultrasound imaging generally requires contact of tissue through a mechanical index matching
medium.
Most clinical diagnostic ultrasound systems operate in the range of 2 to 10 MHz, with the
lower portion of the range used when increased depth penetration is necessary. In most large
patients, a frequency of 3.5 MHz is satisfactory while 5 or 7.5 MHz can often be used in thin
patients or children. These frequencies typically provide depth penetration on the order of 10-20
cm with axial resolutions of 200 - 400 um. Lateral resolutions for clinical systems are around 2
- 4 mm [8]. These parameters are not sufficient for imaging cellular features, but they provide
sufficient resolution and depth penetration for essential diagnostic applications in several
disciplines. Real-time cardiac ultrasound is a standard technique for assessing indicators of
contractile and valve function such as myocardial wall thickness and ejection fraction [9].
Ultrasound has also been applied extensively for monitoring fetal development in utero and for
guidance of minimally invasive surgical procedures. Furthermore, endoscopic ultrasound is
recognized as a potentially important tool for the diagnosis and staging of esophageal, gastric,
colorectal, pancreatic, and biliary tumors [10].
High-frequency ultrasound between 20 and 100 MHz has offered improved imaging of tissue
and cellular microstructure. At 30 MHz, axial and lateral resolutions of about 60 um and 250 um
respectively can be achieved, while extension to 100 MHz provides resolutions down to 19 um
axial and 60 um lateral. Penetration depths for these systems are limited to 4-6 mm depending
on the tissue type [11]. Clinical applications to date have included ocular, skin, intravascular,
gastrointestinal, and cartilage imaging. In addition, high-frequency ultrasound has been
suggested as a tool for experimental work in developmental biology and in tumor biology [12].
Systems for imaging the anterior segment of the eye and for intravascular imaging (IVUS) have
shown particular promise and commercially available models are gaining acceptance in clinical
practice [13, 14]. Despite some successes of high-frequency ultrasound, resolution and contrast
are to date insufficient for imaging of cellular structure in dense tissues. As such, the goal of
10
creating a real-time ultrasound biopsy tool to supplement optical histology for early disease
diagnosis remains largely unfulfilled.
1.2.2
Optical Coherence Tomography
Optical Coherence Tomography (OCT) is a recently developed imaging modality that uses
broadband light sources and low-coherence interferometry to generate high-resolution cross
sectional images of tissue microstructure [15]. OCT generates images by mapping optical
backscatter as a function of depth and transverse position. Because the propagation speed of
light is much faster than photodetector response times, pulse echo time delays cannot be
measured electronically as in ultrasound. To measure backscatter, OCT systems instead use a
fundamentally different technique based on a device called a Michelson interferometer to extract
time delays. Figure 1.2 illustrates the principle of low-coherence interferometry with a
Michelson interferometer. Light from a source is divided into a scanning reference path and a
sample path. The backscattered light probing the sample is recombined with the reference path
light at a photodetector to produce interference fringes. If the light source is monochromatic,
interference is seen over a wide range of reference arm path lengths. If a broadband light source
is used, however, interference will only be seen when the reference arm path matches the sample
path to within the coherence length of the light source. This coherence length determines the
size of the sample volume probed and hence the axial resolution of the OCT system. The
coherence length of the light source varies inversely with the bandwidth of the source.
Increasing the wavelength range of the source therefore reduces the duration of the coherence
gate and provides increased axial sectioning capability [16].
Am
Reference
BS
Sample
Source
Long Coherence Length
Alc
AM
Detector
Short Coherence Length
Alc < Am
Figure 1.1. Schematic illustrating the principle of coherence gating. A Michelson
interferometer is used to combine light from the sample with light passing through a
scanning reference path. For broadband light sources, interference is seen only when the
reference path length matches the sample path length to within the coherence length of the
light source.
11
Scanning the reference arm path length and plotting the envelope of the interference as a
function of this path length generates a map of the backscattered light intensity from the sample.
To generate two-dimensional images, the sample is translated with respect to the incident beam
or the incident beam is scanned across the sample. Typical transverse image dimensions are 3-4
mm. A schematic describing the generation of an OCT image is provided in Figure 1.2.
Standard clinical OCT uses a superluminescent diode laser source and provides cross-sectional
imaging with 10-15 um axial and transverse resolution [17]. High-resolution OCT using modelocked lasers can achieve 1-2 um axial resolution and 5-10 um transverse resolution [18-21].
Transverse Scanning
Backscattered Intensity
Axial
Position
(Depth)
2D Grey Scale or False Color
Image of Optical Backscattering
Figure 1.2. Description of the formation of an OCT image. The backscattered intensity is
mapped as a function of depth. A two-dimensional image is formed by translating the
incident beam with respect to the sample or vice versa.
Contrast in OCT is generated by inhomogeneities in tissue scattering properties and changes
in refractive index. As in ultrasound, there exists a tradeoff between resolution and penetration
depth in OCT images. Higher frequency (shorter wavelength) optical radiation enables
improved resolution at the cost of lower penetration. Use of near-infrared wavelengths between
800nm and 1300nm has enabled OCT image penetration depths of 2-3 mm [17]. In addition to
source wavelength, the penetration depth in OCT images depends on system sensitivity and
incident power. Interferometric detection is an implementation of optical heterodyne detection,
whereby the electric field of a very weak reflection from the sample is measured through
comparison with a strong reference field. Typical system sensitivities to reflected signals of -90
to -100 dB can be achieved. With incident power for in vivo imaging in highly-scattering
tissues restricted by laser safety exposure limits to between 5 -20 mW, signals around 101 W
are respresented in OCT images [22].
Commercially available fiber optic components from the telecommunications industry have
provided a strong base of technology for OCT systems. Development of specific OCT
technology has focused on broadband light sources, high speed scanners, and novel delivery
devices [18-21, 23-29]. The creation of real-time imaging systems and non-contact, minimally
12
invasive imaging probes has enabled investigation of a number of clinical applications in
ophthalmology [30-35], cardiology [36-40], gastroenterology [41-47], urology [48, 49], and
dermatology [48-51]. As in high-frequency ultrasound, the greatest successes to date have been
in ophthalmology and cardiology. Commercialized OCT systems have been tested at several
sites and there appears to be well-defined diagnostic indications for the technology. In other
highly scattering tissues, however, the task of identifying and grading early stage disease has
been more difficult. Assessment of early dysplastic changes with OCT remains an important and
open challenge. While there is some indication that this can be achieved at the level of tissue
morphology, in many cases it appears that cellular-level diagnostics are required. Cellular
imaging with ultrahigh-resolution OCT has been demonstrated in the semi-transparent tissues of
the Xenopus laevis tadpole [52], but clinical imaging of cellular structure in highly scattering
human tissues with OCT has not yet been achieved.
1.2.3
Confocal Microscopy
Confocal microscopy was first proposed by Marvin Minsky in the late 1950s and patented by
him in 1961 [53]. Inspired by a frustrating experience imaging densely packed neurons during
his doctoral thesis work, Minsky sought a technique that could collect light from each individual
point of the specimen, ignoring unwanted scattered light [54]. His elegant solution was a
microscope that uses pinhole apertures to block unwanted light from the detector. Figure 1.3
illustrates the basic principle of confocal microscopy in reflection geometry [55-57]. A point
source illuminates a sample plane through a focusing objective lens. The in-plane backscattered
light is recollected by the objective lens and focused through the point detector. Unwanted
scattered light from outside the focal plane is also recollected by the objective, but this light is
defocused at the detector and is therefore minimally detected. The spatial discrimination against
out of focus scattered light is known as confocal gating.
Focal Plane
Object out
of focus
Scattering
Object
i
L
-
-
-
L2
---
P
Figure 1.3. Illustration of the principle of confocal microscopy. Confocal detection allows
rejection of scattered light from out of the focal plane because this light is defocused at the
point detector.
13
The combination of focused illumination and spatially filtered detection reduces blurring,
increases effective resolution, and improves contrast through improved signal to noise ratio [58].
The transverse resolution varies inversely with the numerical aperture (NA) of the objective lens,
and the axial sectioning capability of the confocal microscope varies inversely with the square of
the NA. Hence, image quality in scattering objects depends strongly on the use of high
magnification, high-numerical aperture objectives. With such lenses, confocal systems can
achieve 1-3 urn axial sectioning capability and better than 1 um transverse resolution.
To generate an image in two dimensions, several scanning approaches have been
demonstrated, including sample scanning, objective scanning, and beam scanning [56]. The use
of laser sources marked a major development in confocal microscopy [59, 60] and enabled high
speed, high resolution point scanning systems at multiple wavelengths. In contrast to OCT, the
confocal laser scanning microscope (CLSM) samples an en face scan plane. Figure 1.4
compares cross-sectional and enface imaging planes.
Transverse
X
Transverse
X
DEPTH PRIORITY
EN FACE
Incident
Incident
Beam
Beam
-~--
~~~~~
-
z Axial
~~
(Range Depth)
y
z Axia l
~~
(Range Depth)
,
y
Figure 1.4. Schematic comparing scanning modalities for Optical Coherence Tomography
(OCT) and Confocal Laser Scanning Microscopy (CLSM). OCT uses depth scanning to
form cross-sectional images. CLSM uses transverse scanning to form enface images.
A second classification of scanning confocal microsopes is known as the Tandem Scanning
Microscope (TSM) [61] and its development has proceeded in parallel with laser point scanning
systems. TSM systems typically use arc lamp light sources and provide real-time, direct view
imaging by scanning the object and image planes in tandem through a perforated rotating disk
known as a Nipkow disk. They offer high-speed scan rates but generally suffer from poor light
efficiency as well as mechanical and optical complexity.
In vivo imaging of unstained tissues using confocal microscopy was first demonstrated using
tandem scanning systems in the cornea of a frog [62]. Extension of TSM systems to human skin
provided exciting, high-resolution images of cellular features at varying depths through the
epidermis [63, 64]. Shortly after, the confocal laser scanning microscope (CLSM) was
demonstrated for in vivo cellular imaging of human tissues [65]. Advances in instrumentation
and design led to the development of video-rate CLSM systems capable of reliable imaging in
clinical applications [65-67]. These systems offer high power illumination and extension to
deeper penetrating wavelengths in the near infrared. Operating at wavelengths of 800 nm and
1064 nm, the systems provide lateral resolution of 0.5 - 1 um and axial sectioning capacity of 3 5 um. Results of CSLM imaging of human skin [67-71] and oral mucosa [72] have
demonstrated capability to explore normal and pathologic cellular features in vivo with
14
impressive correlation of confocal images with histology. Commercial versions of the CSLM
imaging system now offer new tools for clinical diagnostic applications in dermatology and other
specialties where open access to tissue specimens is possible.
In principle, confocal microscopy uses focused illumination and spatially filtered detection to
isolate the single-backscattered component of reflected light from tissue. The image penetration
depth is limited by the ratio of signal to background. Background is determined by the amount
of light entering the finite-sized pinhole from outside the focal volume and the amount of
multiply scattered light that is channeled into the collection volume. In practice, this background
level presents severe limits on the achievable imaging depth and contrast in highly scattering
tissue [73]. Confocal sectioning is weaker than the exponential scattering character observed in
tissue, and the isolation of the single-backscattered component is therefore quickly outstripped
by extinction of the incident light. Moreover, unlike OCT which isolates reflected light
according to path length, confocal microscopy has no intrinsic way to remove multiple-scattered
light from the detected signal. As a result, the most optimized CSLM systems operating at
maximum safe exposure levels have been limited to imaging depths below 300 - 500 um in
human skin and oral mucosa [66]. This penetration depth is sufficient only for imaging of
epithelial layers in certain areas of the skin and gastrointestinal tract, and limits the applicability
of CSLM systems for widespread in vivo clinical application.
The utility of CSLM systems for clinical application is also currently limited by a lack of
miniaturized probe technology. The requirement of high quality, high numerical aperture
objectives makes miniaturization difficult. Typical CSLM systems use bulky, multi-element
lenses and 2-3 times overfilling of the lenses to achieve adequate sectioning. Design of
equivalent optical systems smaller than the 3-5 mm diameter endoscope ports is a daunting task
that limits CSLM imaging to primarily surface tissues. Development of fiber optic CSLM
systems, miniaturized lenses, and micromechanical scanning technology has progressed in recent
years and promises to make CSLM a more complete tool for clinical applications in the future
[74-78].
To improve upon the imaging depth and contrast of confocal microscopy in highly scattering
tissue and to relax design criteria for miniaturized probes, the investigation of alternate,
enhanced sectioning techniques is of prime importance in optical diagnostics research. Two
photon microscopy and optical coherence microscopy are two such techniques that offer unique
improvements over confocal microscopy for optical biopsy.
1.2.4
Two-Photon Microscopy
The physical principal behind two-photon excitation microscopy is the simultaneous
absorption of two infrared photons by a chromophore that induces an electronic transition
normally requiring an ultraviolet photon. The energy transition is bridged by two photons of half
the gap energy rather than a single photon of adequate energy. The theoretical foundation for the
effect was described by Maria Goppert-Mayer in 1931 and was applied for high-resolution
microscopy by Denk and Webb in 1990 [79]. As in confocal microscopy, two-photon
microscopy illuminates tissue with infrared light focused through high-numerical aperture
microscope objectives to micron spots in tissue. The infrared light is absorbed through a twophoton process by endogenous fluorophores in the tissue, which then reemit the energy as
incoherent fluorescence light. The fluorescence is collected by the illuminating objective and
15
spectrally separated from the longer wavelength excitation light before being detected with a
photomultiplier tube. An image is formed by raster scanning either the sample or the beam as is
done in confocal microscopes. Typical two photon systems for tissue imaging use Titanium
Sapphire mode-locked lasers to produce excitation wavelengths between 700 - 900 nm and
collect fluorescence emission in the range of 400 - 600 nm contributed by abundant
biomolecules such as NADH, NADPH, and flavoproteins. Video rate two-photon microscopes
similar to CSLM systems have been realized [80].
The two-photon excitation probability is significantly less than the one-photon probability
and appreciable two-photon absorption occurs only at the focal point, a region of high temporal
and spatial concentration of photons. High spatial concentration results from high numerical
aperture focusing into the tissue. High temporal concentration of photons is achieved using high
peak power mode-locked lasers. Due to the precisely localized two-photon effect in the tissue,
pinholes are not required for spatially filtered detection as in confocal microscopy. Substantial
fluorescence occurs only from the focal volume and can be detected uniquely at the fluorescence
emission wavelength. Two-photon fluorescence intensity depends quadratically upon the
excitation photon flux, which decreases rapidly away from the focal plane.
The precise depth discrimination provided by two-photon microscopy enables powerful 3D
cellular imaging capability to depths not possible with standard confocal microscopy [81].
Preliminary imaging of mouse skin ex vivo and human skin in vivo provides some of the highest
quality cellular images of unstained skin to date [82, 83]. Additionally, the two-photon
technique is intrinsically sensitive to biochemical information since the fluorescence emission
bandwidth can be resolved into contributions from various fluorophores. This opens up
possibility for functional as well as structural imaging at the cellular level.
Unfortunately, two-photon microscopy has limitations that make it practically difficult to
implement for clinical applications. First, as with confocal microscopy, the optical design
requirements are severe. High numerical aperture objectives are needed to produce tiny focal
volumes. Additionally, the need for delivering femtosecond laser pulses to the tissue prevents
use of fiber optics, which will disperse the pulse and reduce peak power. These constraints make
development of miniaturized probes difficult. Second, the high peak-power of femtosecond laser
pulses, while enhancing the two-photon effect, also presents a potential for photo-induced tissue
damage from one photon absorption in the tissue volume. Even the levels used for deep tissue
imaging in previously published in vivo human skin studies are possibly beyond maximum
permissible exposure limits.
1.3
Optical Coherence Microscopy
Optical coherence microscopy (OCM) is a technique which combines the coherence-gated,
heterodyne detection of optical coherence tomography (OCT) with the high transverse-resolution
of confocal microscopy. Figure 1.5 compares the focusing regimes of optical coherence
tomography and optical coherence microscopy. Optical coherence tomography systems typically
use relatively low numerical aperture focusing optics in order to preserve depth of field over the
length of the image depth scan. Optical coherence microscopy by contrast uses high numerical
aperture optics to provide small focal spots. Because the depth of field is severely restricted
when focusing tightly, OCM generally scans an en face imaging plane similar to confocal
microscopy.
16
Low NA
High NA
z
b
+
4--dx
b
-
dx
Figure 1.5. Illustration of focusing characteristics used for OCT and OCM. OCT systems
operate with a relatively low numerical aperture to preserve depth of field over the range of
the image depth. OCM systems use high numerical aperture optics to provide small focal
spot sizes. Note that the axial coherence gating is set by the characteristics of the light
source and is independent of the focusing optics.
OCM systems are typically modified OCT systems.
They consist of a Michelson
interferometer with a confocal microscope in the sample arm. Rather than scanning path length
in the reference arm, the path length is set to match the distance to the focus of the sample arm
and the reference light is phase modulated to provide an oscillating interference signal at the
detector. The systems offer the advantage of highly specific and sensitive optical heterodyne
detection together with nearly an order of magnitude improvement in lateral resolution over
conventional OCT systems.
The heterodyne detection process alone, regardless of the low coherence gating effect,
provides enhanced detection of small signals and improved rejection of out of plane scattered
light as compared to confocal microscopy [84]. Heterodyne detection is sensitive to the
amplitude rather than the intensity of the reflected light, and it provides optical amplification via
a reference arm signal to effectively increase the detectable level of small reflections.
Furthermore, heterodyne detection is phase sensitive detection. Since unwanted scattered light
within the confocal window loses coherence, there is some degree of rejection of this light from
the detection process itself.
Incorporation of broadband light sources in a heterodyne microscope provides path length
gating. The combination of spatial discrimination from the confocal gate with path length
discrimination from the coherence gate provides improved axial sectioning against unwanted
scattered light from outside of the focal plane [85]. The effective axial point spread function
from coherence gating depends on the source spectral shape. For a typical Gaussian spectrum
used in OCT, the axial PSF will also approach a Gaussian function of distance from the object
plane. This sectioning character is stronger than the exponential extinction of incident light and
is stronger than the spatial discrimination provided by the confocal gate alone. Additionally,
coherence gating provides path length selectivity against image-degrading multiply scattered
17
light that is not provided by confocal gating [73]. Figure 1.6 demonstrates the improved axial
sectioning of a coherence gated microscope. These enhanced sectioning qualities can yield
improved contrast and greater imaging depths than standard confocal microscopy [85].
Unlike confocal microscopy, in optical coherence microscopy the axial sectioning can be
separated from the transverse resolution. Using broadband laser sources, the coherence gate can
be set independently of the focusing optics. This principal may be used to relax design
constraints on probe technology for clinical applications. In confocal microscopy, the axial
sectioning is critically dependent on high quality, high numerical aperture focusing optics.
0~
-10-000.-40CD
Confocal
-0
-50-
Confocal +
FDnc -70-
Coherence Gate
'-80-
-90-
,,,
-100 -ii
-200
-100
100
0
Distance (pmn)
200
Figure 1.6. Demonstration of improvement in axial sectioning with combined confocal and
coherence gating compared with confocal gating alone. Enhanced rejection of out of plane
scattered light can improve image contrast and imaging depth achievable with confocal
microscopy alone. Plots reproduced with permission from reference [85].
The axial sectioning ability degrades as the inverse of the square of the numerical aperture. The
transverse resolution, however, degrades only as the inverse of the numerical aperture, not its
square [55]. Because of the weaker loss of transverse resolution with reduced numerical
aperture, there may exist a region of numerical aperture where the transverse resolution is
sufficient for cellular imaging but the axial resolution is not. Combining coherence gated
sectioning from ultra-broadband light sources can provide the necessary axial resolution
independent of the probe optics and may therefore enable cellular imaging with fiber optic probe
designs that would be insufficient for confocal microscopy alone.
Figure 1.7 illustrates the anticipated niche for optical coherence microscopy among other
clinically available imaging modalities. All techniques displayed suffer from generally poor
intrinsic signal contrast in tissue when compared to histologic analysis of stained tissues. With
no other currently available methods, however, these techniques offer the best short term hope
for in vivo, real time assessment of tissue microstructure. OCM has potential to extend and
improve the high-resolution capability of confocal microscopy to depths approaching 1 mm or
more. A reliable tool for visualizing cellular structure at such depths could find applications in
cancer diagnostics and surgical pathology as well as in basic in vivo investigations of such
processes as inflammation and wound healing. In cancer progression, for example, the invasion
of malignant epithelial cells through the basement membrane separating the epithelium from the
18
underlying stroma and connective tissue is an important prognostic finding. The ability to assess
the presence and extent of invasion in vivo and in real time would be an important advance in
early disease detection and staging. Furthermore, in organs such as kidney and spleen, a tough
fibrous capsule surrounds the parenchyma. Imaging through this capsule is difficult with
confocal techniques and may be extended with coherence imaging modalities.
1 mm
Standard
Clinical
0
ULTRASOUN
Z
0
High
Frequency
0
OPTICAL COHERENCE
1 ptm 'OPTICAL
TOMOGRAPHY
A
COHERENCE MICROSCOPY
CONFOCAL MICROSCOPY
100 pm
1 mm
1 cm
10 cm
IMAGE PENETRATION (log)
Figure 1.7. Schematic comparison of high resolution techniques for imaging tissue
microstructure. Optical coherence microscopy has potential for extending the high
transverse resolution of confocal microscopy to depths approaching that of optical
coherence tomography.
1.4
Previous Work on OCM
The concept of optical heterodyne imaging was proposed by Korpel and Whitman in 1963
[86] and the optical heterodyne scanning microscope was then demonstrated in 1973 by Sawatari
[87]. Sawatari's bulk optical system used a continuous wave Helium-Neon laser source for
illumination of the sample and an acoustic beam deflector operating at 70 MHz to produce a beat
frequency in the interference. Optical heterodyne imaging requires phase coherent wavefront
alignment between reference and sample arm beams at the detector, which translates into
directional selectivity similar to confocal imaging. The heterodyned confocal microscope
provides near shot noise limited detection, rejection of incoherent background light, and access
to phase information about the sample.
Optical heterodyne reflectometry with low coherence light was first demonstrated for fault
measurement in fiber optics [88, 89] and applied shortly after for measurement and imaging in
the eye [15, 90, 91]. Izatt et. al introduced the combination of low coherence interferometry with
the optical heterodyne scanning microscope as optical coherence microscopy in 1994 [85].
Izatt's system used a 20x, 0.4 NA objective lens, a fiber optic Michelson interferometer and a
19
superluminescent diode with 30 nm bandwidth centered at 830 nm. This setup provided a
coherence gate of 18um and a confocal gate of 22 um. Phase modulation in the reference arm
was performed with a fiber stretching piezo-electric device, which produced less than 1 um of
path length variation. Images were generated by raster scanning a sample under the microscope
with slow scanning stages. Imaging results of a polymer microsphere suspension were used to
verify a single scattering model and to demonstrate potential for imaging up to several hundred
micrometers deep or between 2-3 times the depth of standard confocal microscopy. Kempe and
Rudolph demonstrated similar enhancement of axial sectioning and image contrast in a
microsphere scattering model using a bulk interferometer system and Ti:Sapphire solid state
laser source [84, 92].
Izatt applied his system design for OCM imaging in human gastrointestinal tissue at 1300 nm
[93]. A superluminescent diode with 47 nm bandwidth centered at 1299 nm provided a
coherence gate of 15.9 um. The confocal gate and transverse resolution using a 40X, 0.65 NA
objective were 5 um and 1.9 um respectively. Phase modulation was performed with a
piezoelectric stack at 1.64 kHz. Better than 95 dB sensitivity was achieved with 140 uW on the
sample. The system allowed visualization of epithelial cells at depths greater than 500 um in the
colonic mucosa, clearly demonstrating range of penetration in tissue superior to confocal
microscopy alone.
Schmitt demonstrated a novel technique for generating high-resolution OCT cross-sectional
images in human skin by scanning the reference arm together with the sample arm on a slow
scanning stage [94]. The technique helped to compensate the relative slip of the confocal and
coherence gates that occurs when focusing deep into tissue and removed the depth of field
limitation encountered in standard OCT scanning modes. Lexer et al. extended this concept of
focus-tracking to higher speeds with demonstration of a dynamic coherent focus method
whereby a galvanometer mirror in the sample illumination path was used to scan the focus depth
in the sample [95]. This technique was demonstrated for moderate speed of 1 image per second
with a transverse resolution of 5 um. The setup uses a bulk interferometer and three scanning
mirrors in the sample path, making design of fiber optic miniaturized probes with this technique
unlikely. Broadband operation of this system has also not been demonstrated.
Several approaches have been pursued for the development of fast-scanning en face OCM
systems. Podoleanu et al. used a Newton rings sampling function to acquire en face images of
the human retina [96-98]. Spatial resolution of 6 um and image acquisition rates up to several
frames per second. Performance equivalent to confocal microscopes has not been demonstrated
with this system design. Furthermore, the technique decodes images based on quasimonochromatic light source assumptions, making its utility for broadband coherence imaging
uncertain.
Beaurepaire et al. demonstrated the principal of full-field optical coherence microscopy using
a parallel detection technique [99]. With a spatially incoherent source, speckle-free images with
diffraction limited resolution were acquired without scanning. A special Michelson objective
lens illuminated the sample and a photoelastic modulator provided path difference modulation.
The system suffers from complex optical design requirements and has not been demonstrated for
use in fiber optic delivery devices.
Westphal et al. demonstrated a fast point-scanning OCM system similar to commercial
confocal microscope designs for cellular imaging in human skin [100]. En face images were
demonstrated with 5 um axial sectioning and better than 2 um lateral resolution to depths of 600
um. The setup used a high power broadband superluminescent diode laser source centered at
20
1310 nm with 67 nm bandwidth, providing a coherence gate of about 12 um. A commercial
electro-optic phase modulator provided the heterodyne beat signal for detection and fast resonant
galvanometer scanners were used to achieve up to 8 frames per second imaging capability. The
system suffered from relatively low system sensitivity of 76 dB and bandwidth limitations
imposed by the phase modulator.
1.5
Scope of Thesis
Development of high-resolution, high-speed OCM imaging systems is an important area of
research toward the creation of an optical biopsy tool for in vivo clinical applications. To take
full advantage of the improved axial sectioning provided by coherence gating, OCM systems
should be designed to support large optical bandwidths available with femtosecond laser sources.
Construction of real-time, broadband OCM imaging systems has previously been limited by the
availability of high-speed, broadband phase modulators. Earlier work has used either a fiberstretching piezoelectric modulator, which limits speed, or a waveguide electro-optic phase
modulator, which limits the optical bandwidth of the system. Furthermore, waveguide devices
are commercially available only at select wavelengths. This thesis discusses the demonstration
of a novel, broadband OCM system that enables real time imaging of cellular structure in highly
scattering tissue. The system integrates a high-resolution OCT system with a reflective grating
phase delay modulator and a fast scanning confocal microscope. Grating phase delay scanners
have been developed and demonstrated previously for high speed OCT imaging and for phase
modulation [101, 102]. The novel reflective geometry demonstrated here enables OCM imaging
with large bandwidth, providing coherence gates of only a few micrometers. Moreover, the
flexible OCM system design can readily be implemented at wavelengths that were previously
inaccessible for OCM imaging.
The broadband OCM system is used to demonstrate a new operating regime for real-time, in
vivo imaging of cellular structure in human tissues. Combined coherence and confocal gating is
shown to relax microscope design constraints imposed by confocal microscopy. In particular, a
short coherence gate is used to enhance weak confocal sectioning, thereby enabling cellular
imaging in situations when confocal microscopy alone would be inadequate. The results
demonstrated offer promise for cellular imaging in clinical applications that require probe
technology unsuitable for confocal imaging.
The remainder of this thesis is divided into four chapters. Chapter 2 describes the underlying
principles of optical coherence microscopy. Analyses of confocal microscopy and low
coherence interferometry are individually presented and then combined to describe OCM image
formation. The effects of combined confocal and coherence gating are discussed to develop an
understanding of OCM operating regimes for in vivo imaging. In addition, the theory of
operation of the grating phase modulator is provided.
Chapter 3 discusses development and characterization of the broadband OCM system.
Design constraints for in vivo imaging are described followed by individual discussions and
measurements of the system components. Particular attention is devoted to design and
characterization of the reflective grating phase modulator and the sample arm microscopes used
for tissue imaging.
Chapter 4 demonstrates the OCM system for high resolution, in vivo imaging in an animal
model and in human skin. High-resolution cellular images of Xenopus laevis tadpole and human
21
skin are presented and discussed. Particular emphasis is placed on demonstration of the use of a
short coherence gate to enhance weak confocal sectioning.
Finally, chapter 5 summarizes the results and discusses future work.
22
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Youngquist, R., S. Carr, and D. Davies, Optical coherence-domain reflectometry: a new
optical evaluation technique. Optics Letters, 1987. 12(3): p. 158.
Fercher, A.F., K. Mengedoht, and W. Werner, Eye-length measurement by interferometry
with partiallycoherent light. Opt Lett, 1988. 13: p. 1867-1869.
Swanson, E.A., et al., High-speed optical coherence domain reflectometry. Optics
Letters, 1992. 17: p. 151-153.
Kempe, M. and W. Rudolph, Scanning microscopy through thick layers based on linear
correlation. Optics Letters, 1994. 19(23): p. 1919-1921.
Izatt, J.A., et al., Optical coherence tomography and microscopy in gastrointestinal
tissues. IEEE Journal of Selected Topics in Quantum Electronics, 1996. 2(4): p. 1017-28.
Schmitt, J.M., S.L. Lee, and K.M. Yung, An optical coherence microscope with enhanced
resolvingpower in thick tissue. Optics Communications, 1999. in press.
Lexer, F., et al., Dynamic coherentfocus OCT with depth-independent transversal
resolution. Journal of Modern Optics, 1999. 46(3): p. 541-553.
Podoleanu, A., et al., Coherence imaging by use of a Newton rings samplingfunction.
Optics Letters, 1996. 21(21): p. 1789-1791.
Podoleanu, A.G., et al., Simultaneous en-face imaging of two layers in the human retina
by low-coherence reflectometry. Optics Letters, 1997. 22(13): p. 1039-1041.
Podoleanu, A.G., G.M. Dobre, and D.A. Jackson, En-face coherence imaging using
galvanometer scanner modulation. Optics Letters, 1998. 23(3): p. 147-149.
Beaurepaire, E., et al., Full-field optical coherence microscopy. Optics Letters, 1998.
23(4): p. 244-246.
Westphal, V., H.W. Wang, and J.A. Izatt. Real-time in vivo optical coherence
microscopy. in Conference on Lasers and Electro-Optics (CLEO). 2001. Baltimore, MD.
Tearney, G.J., B.E. Bouma, and J.G. Fujimoto, High-speedphase- and group-delay
scanningwith a grating-basedphase control delay line. Optics Letters, 1997. 22(23): p.
1811-1813.
Zvyagin, A.V. and D.D. Sampson, Achromatic opticalphaseshifter-modulator.Optics
Letters, 2001. 26: p. 187-190.
27
28
Chapter 2
Optical Coherence Microscopy
2.1
Overview
Optical coherence microscopy (OCM) combines high sensitivity, coherence-gated optical
heterodyne detection with the high transverse resolution and spatial discrimination against out of
focus scattered light provided by a confocal microscope. Figure 2.1 shows a block diagram of
the OCM system discussed and demonstrated in this thesis. Low coherence light is split equally
into reference and sample arm paths by a fiber optic coupler. The reference arm light passes
through a delay modulator while the sample is illuminated through a scanning confocal
microscope.
Backreflected
light
from the two
arms is recombined
at dual balanced
photodetectors to produce a heterodyne interference signal, which is then amplified, filtered, and
demodulated. The demodulated signal is digitized and displayed on a computer screen. Dual
balanced detection is employed to eliminate excess common mode laser noise on the reference
and sample arms. Polarization controllers in both sample and reference arms ensure electric field
alignment for maximum interference.
Broadband Light Source
Grating Phase
Delay Modulator
Polarization
Control
50/50
D1
Detection
50/50
Electronics
Polarization
Control
I
D2
Computer
-S-VHS
Moitor
Recorder
IScanning
Function
Generator
Galvo
F n Controllers
Figure 2.1.
imaging.
Confocal
*Microscope
X
d
-00'modulator
Schematic of broadband optical coherence microscopy system for in vivo
This chapter provides background theory of operation of the OCM system and discusses
important design criteria and principles of OCM image formation in biological tissues. Section
2.2 relates the parameters used to describe the optical properties of tissue and provides a simple
29
picture of the scattering processes that generate the image signal. Sections 2.3 and 2.4 discuss
important principles of confocal microscopy and low coherence interferometry and section 2.5
combines these principles to describe image formation and operating regimes for OCM. Finally,
section 2.6 discusses operation of the grating phase delay line used in the OCM system to
generate the heterodyne signal.
2.2
Scattering in Biological Tissues
Biological tissues consist of a network of cells, vessels, and other structures suspended in a
mesh of collagen and elastin fibers. For optical imaging, this translates to a sample with
turbulent refractive index variations that distort the spatial and temporal coherence of the sample
beam [1]. A number of different scattering processes are at work in dense biological tissues.
These are illustrated schematically in figure 2.2, excerpted from published work by Schmitt [2].
Source
Reference
Detector
E,' (gt)
Er (PLt)
f(Fr *(Pt+,I) E,' pt) d2p
(
Single backscatter
Wide-angle scatter
00
Phase-front distortion by
0
large-scale index variations
Low-angle multiple scatter
0
00
0
00
Illustration of important scattering processes in dense biological tissues.
Figure 2.2.
Excerpted from reference [2]
For imaging microstructure, it is desirable to isolate the single backscattered component of
light returning from tissue. This component, however, is obscured by multiply scattered light
arising from above and below the focal plane. Axial sectioning techniques provided by confocal
microscopy and by low-coherence interferometry offer ways of reducing the amount of multiply
scattered light that contributes to the image signal.
Because of the complex heterogeneity in tissue, full application of electromagnetic theory to
analytically describe propagation and scattering is impractical. Simple models of biological
tissues treat them as suspensions of discrete objects for which the analytical scattering solution is
available. Typically, spherical scatterers are assumed so that Mie theory can be applied [3-5].
Numerical simulations based on Monte Carlo and finite difference time domain (FDTD) methods
have also been used to analyze scattering from more complex models of cells and their
environment [6, 7]. The preferred description of light propagation in biological tissue, however,
30
abandons analytic approaches in favor of radiative transport theory [8, 9]. Transport theory
allows a simple description of propagation based on absorption and scattering coefficients that
define energy loss through the medium.
The absorption and scattering coefficients, p, and p, respectively, are grouped into a total
total attenuation coefficient given by
P, = Pa + Ps
(2.1)
The transmittance T of tissue can then be described as an exponentially decaying function of
depth into the medium, where the rate of decay is set by the total attenuation coefficient. This is
typically known as Beer's law.
T = e-"a
(2.2)
The absorption and scattering coefficients are in units of inverse meters m-1 . They can be used to
define a normalized unit of depth called the transport mean free path 1, given by
it = 1
(2.3)
Pt
The near infrared (NIR) wavelength range of 600 nm to 1300 nm is known as the
"therapeutic window" because absorption is relatively smaller than in the ultraviolet or the far
infrared. Weak absorption allows deeper penetration into tissue, leading to greater imaging
depths. For this reason in vivo imaging techniques are designed for operation with NIR
wavelengths. Furthermore, scattering itself varies inversely with wavelength meaning that the
far edge of the therapeutic window provides optimum penetration depth. The penetration at
1300 nm is more than two times greater than the penetration at 800 nm. Most optical coherence
tomography systems for deep tissue imaging therefore use wavelengths around 1300 nm [10].
31
2.3
Confocal Microscopy
Confocal microscopes provide sectioning ability in scattering media. A point source
illuminates a sample plane through a focusing objective lens. The in-plane backscattered light is
recollected by the lens and focused through the point detector. Unwanted scattered light from
outside the focal plane is also recollected by the objective, but this light is defocused at the
detector and is therefore minimally detected. The spatial discrimination against out-of-plane
scattered light is known as confocal gating. Figure 2.2 illustrates this principle as it is typically
implemented in a reflection mode confocal laser scanning confocal microscope.
Lens
-- ' '-..
Beam
Splitter...
Source
-.
-.
i
---.
'-.Plane
r.--
Focal
Detector
Figure 2.2. Typical setup for confocal laser scanning microscopes in reflection mode. A
beam splitter divides the optical setup into illumination and detection paths.
This section discusses important background analysis and principles of confocal microscopy
helpful for understanding the role that the sample arm confocal microscope plays in OCM image
formation.
2.3.1
Image Formation in Confocal Microscopes
Application of vector diffraction theory based on Huygen's principle can be applied to obtain
the full electromagnetic field solution at the detector. This approach is tedious and has been
approached elsewhere [11, 12]. Simpler descriptions of the response of the confocal microscope
follow a Fourier optics approach deriving from scalar diffraction theory. We assume a linearly
polarized electric field written as
A
E(r,t) = n V/ (r)ejcot
A
A
(2.4)
A
where ig(r) is a complex number and r = x x + y y + z z is the position vector in three-dimesional
space. As Haus notes, this is not strictly legitimate due to the requirement for the divergence of
the electric field to equal zero in free space [13]. Often this point is ignored for simplicity and
will be done here. The scalar field y(r) satisfies the scalar wave equation
32
(V 2 + k 2 )V/(r)
=
0
(2.5)
where the wave propagation vector k satisfies the dispersion relation given as
k 2 =k
p2
2+kk2
(2.6)
In the paraxial limit, k is inclined by a small angle with respect to the axis of propagation,
assumed to be the z axis here. Following Haus, the paraxial restriction is written [13]
k 2 -k,-k'Y =k-
k=
k+k
x
2k
(2.7)
'
Under this approximation, the scalar field qj(r) can be written as
f(x, y, z)= U(x, y, z)e-kz
(2.8)
with the u(x,y,z) defined by a superposition of plane waves with amplitude distribution
u(x,y,z)
=
jdkjdkUo(k,
~j
kY)
i(k+k)1/2k]z
(2.9)
Setting z = 0, it becomes clear that the function Uo(ks,k,)is the Fourier transform of the
amplitude distribution of the field Vg(x, y, z) at z = 0. Taking the inverse Fourier transform of
(2.9) with respect to x and y at z = 0 one obtains
U0 (k,,)=
dxo
2
f dy0u0 (x0 , y0 )e"(kxxo+kYO)
(2.10)
where (x 0 , yo) are the coordinates in the x-y plane at z = 0. Introduction of (2.10) into (2.9) and
simplification of the integrals yields the Fresnel diffraction integral in the paraxial approximation
.00
C0
f dxO dyOU 0(x0 , y 0 )e
u(xyZ)
Az 00
=
0
+(YYo
)2
/k2z)[(x-xo
)2
(2.11)
h(x, y, z)0uO (x, y)
The field at an arbitrary location (x,y,z) can therefore be expressed as a convolution of the field
at location z = 0 with the Fresnel kernel, defined as
33
h(x,y,z)= J e -jk[(X2.,+>Y2
AZ
(2.12)
1z
By the convolution theorem of Fourier transforms, the distribution can likewise be expressed in
the frequency domain as a product of the transform of the Fresnel kernel with the transform of
the field amplitude distribution at z = 0.
(2.13)
U(k,, k,, z) = (2Zf) 2 H(kx,k,, z)U 0 (k,,k,)
with
(2.14)
12 ej[(k 2+k>/2k]z
H(kx,k,,z)=
(2z)2
The Fresnel kernel defines the action of propagation through a region of free space. A similar
transformation can be defined for other optical elements, such as slabs of material and lenses.
Propagation through a thin lens can be described as multiplication by a complex factor defined as
j
l(x, y) = P(x, y)e 2
2
k (X +Y
2
(2.15)
where P(x, y) is the pupil function of the lens and the complex exponential represents a
parabolic (x,y) - dependent phase delay [14]. The pupil function P(x, y) is in general a complex
function describing the physical extent of the lens, the transmissivity, and any aberrations
introduced. The focal distance f characterizes the lens and is familiar from the lens law of
elementary optics, which relates the object distance do to the image distance d,.
1 = 1+f
do
(2.16)
d,
These results can be applied for analysis of the response of a confocal microscope. Figure 2.3
illustrates the optical geometry for analyzing image formation of u,(x,y) from object field
u (x, y).
u0 (x,y)
u(x,y)
14--
UL(X,Y)
di -1
d10-
Figure 2.3. Schematic geometry for image formation by a lens.
34
u(X,y)
Expressed in convolution form, the image field u, (x, y) can be written
u, (x, y) = h(x, y, d1 ) 0 {l(x, y) [h(x, y, do) 0 u (x, y)]}
(2.17)
where h(x, y, do) and h(x, y, d1 ) represent propagation through distances do and d1 ,
respectively, and l(x, y) represents the effect of the lens. To find the impulse response of the
imaging system, let the object be an impulse 5 function at coordinates (xo, yo). Making use of
(2.12) and (2.15), the field after the lens can now be written as
hL (X
(x2
2
(2.18)
2
jek[(x-x ) +(y-yO) ]12d 0
Ady
Application again of the Fresnel integral (2.11) and simplification using the lens law (2.16)
yields an expression for the impulse response of the imaging configuration expressed in the
coordinates of the image plane (x, y').
k
k
(X2 +2
-i___(X
___X1+~
h,(x,,y1)=-
Here M
-e
Add,-0
o
'
e
2
1 --
2c
Y1 )
c
j-0 1 RX
dx fdyP(x,y)e d
+Mx,,)x+(Y, +Mv 0 )v]
(2.19)
-
is defined as the magnification of the imaging system. The field at the image
do
plane for an arbitrary object field can now be given as a convolution of the object field u (x, y)
with the above impulse response.
=
u, (x, y) = u0 (x, y)0 h, (x, y)
(2.20)
Note that the intensity in the image of a single point object u0 (x, y)
by the square of the magnitude of the impulse response.
I(x, y)= h,(x, y)12
=
3(x)S(y) is given simply
(2.21)
The premultiplying constants and phase variations in the impulse response are important in
certain circumstances but are usually ignored to simplify analysis [14]. In addition, a quadratic,
z-dependent phase factor is added to account for effects of defocus along the axial dimension.
Considering a circularly symmetric pupil of radius a, the impulse response can then be written
in polar coordinates as [15]
,
v 2
h 1 (u,v) = 2JfpdpP(p)J0 (vp)e2'
0
35
2
(2.22)
where J, is a zero order Bessel function of the first kind, p = r / a is the normalized radial
coordinate in the plane of the lens, and (u, v) are normalized optical coordinates related to the
real axial and radial coordinate (z, r, ) around the image plane. The coordinates are defined as
(2.23)
U = - z sin2 (a/2)
and
v=
2z
n
r, sin (a)
(2.24)
where sin(a) is the aperture of the lens.
The results for point illumination through a single lens configuration are readily applied to
the microscope [15]. Figure 2.4 illustrates the unfolded optical geometry for a microscope.
Objective h.
Collector h 2
Detector D
Unfolded optical geometry for a scanning microscope. For a confocal
Figure 2.4.
microscope the detector D becomes very small. In reflection mode, the objective and
collector are identical such that h, = h2.
In scanning optical microscopy, the intensity is typically detected and mapped as a function of
position to create an image. For incoherent imaging with a conventional microscope, the
collector pupil and effective detector are very large, and the image intensity for object
reflectance r is written as
(2.25)
I= hI 2 @ r21
The coherent imaging case is defined by a small collector pupil with a large area detector, and
the image intensity is written
(2.26)
I= h, 9 r12
In the confocal microscope, the detector is replaced with a point detector D(x, y) =(x),(y),
such that the image intensity is always coherent regardless of pupil size. The confocal system is
essentially a coherent imaging system where the point spread function is given by the product of
the point spread functions of the two lenses.
Iof
=
(hh2)0r
2
(2.27)
For reflection geometry, the objective and collector are one and h =hk =h 2 . The three
dimensional image intensity function for coherent imaging of a impulse point object can then be
written
(2.28)
IC,u, v)= h(u, v) 2
36
for the conventional microscope and
(2.29)
Iconoal(UV)= h,(u, v)14
for the confocal microscope.
The field response to an impulse point scatterer in a confocal system is the square of the
impulse response given by (2.22).
hc,,focal(u,v) =[h,(u,v)]
2
(2.30)
The image field for impulse point illumination of a general object is the convolution of the
sample reflectance rs (x, y, z) with the confocal response.
Uconcai
(x, y, z)= rs (x, y, z) 0 [h, (x5 y, z)] 2
(2.31)
For confocal microscopy in scattering media, the responses due to the distribution of scatterers
are integrated at the detector. Ignoring the transverse dependence for simplicity, an axially
distributed collection of scatterers imaged with a confocal microscope produces a DC detector
level described by
ICc Jdz rs(z)9[h,(z)]
(2.32)
where z = 0 corresponds to the position of the focus and rs (z) characterizes the axial
reflectance. Scanning the sample or the beam introduces an (x,y,t) dependence that represents
the en face image intensity.
2.3.2
Lateral Response
Substituting (2.22) into (2.28) and (2.29) and evaluating at the focal plane u = 0 one obtains
the lateral response of the conventional and confocal microscopes for a point object uniformly
illuminated through an ideal pupil defined to be zero for p >1 [15].
j
Icon (0, v) =
Iconfocal (0,v)=
V
(2.33)
(2.34)
The confocal point response offers a sharpened central peak and dramatically reduced
sidelobes, thereby providing enhanced imaging capability. The single point transverse resolution
is defined as the full width at half maximum (FWHM) of the lateral point spread function.
Solution for the half power points of (2.34) above gives the transverse spot size for the confocal
microscope with uniform illumination
0.37Z
dxcofolf,(3dB)
NA
(2.35)
NA
37
where NA = n sin(a) is the numerical aperture of the objective lens [11]. For a Gaussian beam
spot, the transverse spot is more often characterized by the e-2 radius of the lateral response,
which is given as
(2.36)
dxconfoca (e-')=0.6
NA
Gaussian
In practice, measurement of the transverse resolution is often performed by recording the
edge response of the microscope. Assuming point illumination, the 10-90% edge response is
given as
(2.37)
-90%)= 0.44
NA
It can be shown that for a Gaussian beam spot, the e-2 radius is 78% of the 10-90% edge width
and the FWHM is 92.5% of the edge width [16, 17].
The resolution can also be described by the ability to resolve two nearby points, the so called
two-point resolution. The Rayleigh criterion states that two points can be distinguished when
there is a 26.5% drop in intensity between them. The two-point resolution for the confocal
microscope is given as [11]
dxconfcal (10
edge
dofolO,(Rayleigh) = 0.56A
(2.38)
NA
The transverse resolution decreases linearly with increasing wavelength and goes as the inverse
of the numerical aperture of the objective. High numerical aperture and shorter wavelengths
therefore provide increased resolving power in confocal systems.
2.3.3
Axial Response and Sectioning
Substition of (2.22) again into (2.28) and (2.29) and evaluation at v = 0 provides the axial
response of the conventional and confocal microscopes to a point object in the focal plane [15].
ifcolv (U
point
0)-
sin(u 4/ 4)
sin(u / 4)
Iconfoa,(u,0)=
point
2(2.4)
u1
4
.(2.40)
Iu14
Again, the confocal case has lower sidelobe levels and a sharpened central lobe as compared to
the conventional case. The full width at half maximum of the confocal axial point spread
function assuming a uniform point source can be shown to be
dz
(3dB)=
~ 1.24nA
0.62
NA2
n(1-cosa)
where the approximation leads to 2-6 % error for large numerical aperture [18].
38
(2.41)
The optical sectioning ability of a confocal microscope is generally described by its response
to a planar object. For a perfectly reflecting plane at the focus, the response becomes [15]
[sin(u /2)2
confocal (U)
plane
_
u/ 2
(2.42)
_
The full width half maximum span of the response can be shown for uniform point illumination
to be [19]
0.45A
0.90nA
n(1-cosa)
NA2
where again the approximation suffers at large NA.
FWHM of the axial irradiance goes as [16]
(2.43)
For Gaussian beam assumptions, the
dzie (3dB) = 1.2
4
(2.44)
NA
Gauss
Axial sectioning capability decreases linearly with increasing wavelength and goes as the inverse
of the square of the numerical aperture. The square dependence on NA makes axial resolution
critically dependent on the use of high numerical aperture lenses. Figure 2.5 illustrates the
dependence of axial and transverse resolutions on numerical aperture for near-infrared
wavelengths of interest for biomedical imaging. Note the rapid decrease in axial sectioning
capability with lower NA compared to transverse resolution.
20
--- Axial 800nm
-
Transverse 800nm
..... Axial1300nm
Transverse 1300nm
15
E
0
10
5
0
..
0.2
0.4
0.6
-**..
0.8
1
1.2
1.4
NA
Figure 2.5. Comparison of axial and transverse resolution for confocal microscopes using
NIR wavelengths.
39
2.3.4
Effect of Aberrations
Deviations from the idealized conditions of paraxial or Gaussian optics are known as
There are two major types of aberrations, chromatic and monochromatic.
aberrations.
relate to the variation in optical properties of the lenses at different
aberrations
Chromatic
frequencies. Broadband illumination leads to focusing of different wavelengths at different
depths, which broadens the effective axial and lateral responses. The monochromatic aberrations
are also known as Seidel aberrations. They include spherical aberration, coma, and astigmatism,
which degrade the clarity of the image, and Petzval field curvature and distortion, which deform
the image [20]. Spherical aberration is known to be particularly troublesome when imaging
through layered media of varying refractive index at high numerical aperture [21]. An aberrating
layer degrades the axial response, leading to a decrease in peak intensity, a broadening of the
central lobe, and an increase in strength of the sidelobes. Some compensation for spherical
aberration is possible by adjustment of the objective tube length or adjustment of the optical
character of the immersion medium [22, 23].
The effects of aberrations can be incorporated into previous analysis through the objective
pupil function. The aberrations are introduced as phase factors depending on the Seidel
coefficients of spherical aberration A, primary coma B, and primary astigmatism C. In polar
coordinates, the pupil function becomes
.12
P(p,0) = e
ej 2)r(AP 4 +BP3 cOs 0+CP 2 cOs2 o)
(2.45)
where a term dependent on the axial coordinate u has also been included to account for the
effects of defocus [15].
2.3.5
Effect of Finite Detector Size
The analysis thus far has assumed an ideal impulse point detector. A realistic finite-sized
circular detector can be included in the confocal response as
|h,(2u,v)|2 D(v)vdv
I(u) =
(2.46)
where we take h, (u, v) given by (2.22). Choice of pinhole size is a compromise between
maximizing the detected signal and preserving resolution. As the pinhole gets larger, the
microscope loses its confocality and its response approaches that of the conventional
microscope. The depth resolution is less sensitive than axial resolution to pinhole size. Axial
resolution is preserved for a normalized pinhole radius v : 2.5 while lateral resolution is
maintained only to a radius v !0.5. A pinhole size of v, ~1.4 has been determined optimum
for in vivo laser scanning confocal microscopy in scattering tissues [16]. The normalized pinhole
size is defined at the object plane as v,
=
NAr, where r, is the actual pinhole radius at the
A
40
object. The physical pinhole size at the detector is obtained by scaling of r, by the optical
system magnification.
2.3.6
Fiber Optic Confocal Microscopes
The core of a single mode fiber can be used as both the point source and point detector in a
confocal microscope.
These systems differ fundamentally from bulk optical confocal
microscopes [24]. Bulk confocal microscopes behave as partially coherent imaging systems due
to the finite extent of the detector. Fiber confocal systems, however, are intrinsically coherent
imaging systems even for finite values of fiber core size. Reflected light from the sample must
match the fundamental field mode of the fiber in order to couple into it. Compared to (2.46)
where the intensity of light from the sample is integrated over the detector, fiber-optical confocal
systems integrate the field amplitude over the fiber profile and are therefore linear in amplitude.
Ifibe(U)
optic
=
lh,(u,v)f* (v)
vdv
(2.47)
The effective axial point spread function of a planar sample in a fiber optic confocal microscope
is given as [24]
Iplanar
fiber-optic
Here A=
"
(u)-
(1-e
_
-)(Aiu)
(2.48)
describes the normalized fiber spot size with ao the pupil radius of the
objective, ro the radius of the core of the single mode fiber, and d the distance from the fiber tip
to the collimating lens. True confocal detection is preserved when A 1. For typical fiber
system parameters in the near-infrared wavelengths A 1 is typically maintained [25].
2.3.7
Scanning Confocal Microscope Designs
Images in confocal microscopy are generated by either scanning the sample with respect to
the microscope or by raster scanning the beam on the sample at a fixed depth. In either case,
images are created from two-dimensional en face maps of backreflected intensity. For in vivo
applications, the desire for high-speed imaging and access to a range of imaging sites favors
beam scanning techniques over sample scanning. Scanning devices are typically either small
galvanometer-controlled mirrors or rotating scanners.
Linear, waveform controlled
galvanometers offer the most precision and flexibility in scan control, but they are generally
limited in speed to below 1 kHz. Resonant galvanometer scanners offer increased speed to 2
kHz and higher but are not waveform controlled and are limited to sinusoidal scan responses.
Rotating polygonal scanners offer the highest speeds but are bulky and difficult to design into
41
flexible scan probes. Using combinations of these scanners, video rate confocal systems have
been demonstrated [16].
Figures 2.6 and 2.7 illustrate four possible reflection mode scanning confocal microscope
designs based on single mode fiber delivery and detection. Fiber optic implementations offer the
most flexibility for in vivo applications and are considered here. The desired microscope
geometry depends on the type of objective lens used. Finite tube-length objectives create an
intermediate real image of the focal plane at a specified location behind the objective, typically
about 160 mm. Figure 2.6a and 2.6b are suitable designs for finite tube-length objectives.
Infinity corrected objectives create an image of the focal plane at infinity behind the objective,
thereby providing a collimated beam to the intermediate optics. The lens just behind the infinity
corrected objective is often called the tube lens or scan lens. It creates a real image of the focal
plane from the collimated output of the objective and is often precisely designed for
compensation of aberrations produced by the objective. Figure 2.7a and 2.7b illustrate suitable
designs for infinity corrected objective lenses.
The role of the intermediate optics is delivery of a magnified, angle scanning beam to the
objective lens, which then creates a raster scanned image of the fiber core on the sample.
Microscope objectives are designed to be telecentric, meaning that the magnification is
independent of the focus position [11]. As such, light from any point on the sample will pass
through the telecentric plane as a rectilinear beam and a raster scanning beam will appear to
rotate around a point in the telecentric pupil plane. A scanning geometry should be designed to
image the scanning mirrors to this pupil plane. For finite tube length objectives, imaging of the
scanners to the intermediate image plane set by the tube length meets this condition. For infinity
objectives, however, the tube lens must be carefully set to image to the telecentric pupil plane.
In either case, a raster plane in the optical layout can be defined. The field of view (FOV) of the
microscope is then given by the size of the scan in the raster field reduced by the magnification
MO of the objective lens and its tube lens (if needed).
FOV =
dX
XR
(2.49)
MO
Here dXR is defined as the dimension of the scan in the raster plane. The spot size of the
microscope is simply the image of the fiber core at the focal plane of the objective lens. For an
overall system magnification factor M including the objective, the spot size can be written as
.fiber
spot size =
core
M
(2.50)
The chosen geometry of the microscope depends on desired scanner configuration. Closecoupled scanners place X and Y scan mirrors centered about a single plane in the optical layout.
They allow in general for more compact designs with fewer optical elements, which in turn
provides lower loss and higher system throughput. Close-coupled scanners, however, deviate
from ideal imaging geometry since both scanners cannot be simultaneously imaged to the
telecentric pupil plane due to the required physical separation of the mirrors. Separation of the
scanners into distinct image planes allows for an ideal scan configuration at the cost of more
42
optical elements and increased size. Figure 2.6a and 2.7a illustrate separation of the scanners
while 2.6b and 2.7b demonstrate close-coupled scanner configurations.
4
OBJ.
12
3
f
rfiber
*
Lt = Tube Length
(4
So
3
1/Si + 1/So =
(3
d
f2
(2
A1
1/ 14
Si = Lt + 14
1(2
(
OBJ.
Raster
Plane
fiber
X
(1
f2
Lt = Tube Length
d
So
1
Si =Lt + 2
Figure 2.6.
Fiber optic scanning confocal microscope designs for finite tube length
objective lenses.
I
5
OBJ
13
4
f2
n1
Raster
fiber
Telecentric
Pupil Plane,.,
5
(5
B4
(4
3
I
f3
OBJ.
(2
(2
3
fiber
Raster
I
1
i
03
f(
Plane
ln
M
(l
f2
x
T lecentri
Pupil Plane
d
f2
(2
d
n
Figure 2.7.
Fiber optic scanning confocal microscope designs for infinity corrected
objective lenses.
43
2.4
Low Coherence Interferometry
Low coherence interferometry (LCI) using a broadband light source provides path length
gating of light from a sample. LCI provides sensitivity to the field amplitude through optical
heterodyne detection. Interferometric heterodyne detection requires detected light to be phase
coherent with a reference beam over the extent of the detector, and offers amplification of weak
sample arm signals through correlation with the strong reference arm signal. This section
discusses the principles of low coherence interferometry as they apply to optical coherence
microscopy.
2.4.1
Interferometer Analysis
The theory behind low coherence interferometry was presented eloquently by Hee in his
doctoral thesis [26] and will be adapted here. Figure 2.8 shows the basic configuration of a fiber
optic Michelson interferometer with the coupler ports labeled 1-4. The reference and sample
arm field reflectivities are described by complex functions RRej#(iOt) and RS, respectively, where
e-j(o') is a wavelength dependent phase shift introduced by the scanning reference arm. In
general, the fields at the detector are given by the fiber single mode profile. Interference occurs
between fields sharing the same mode profile such that the phase of the reference and sample
arm wavefronts match. The detected signal is determined by integration of the mode profile of
the interfering fields together with the transverse dependence of the detector response. For
simplicity, plane waves of a single linear polarization are considered here and the fields from the
reference and sample arms at the detector are described as ER (co, t) and Es (co, t) .
1
3
Reference R,.eMd*At
Source Eoei'01
E+ ES
Detector D,
ER
XS
Sample Arm Rs
4
2
Figure 2.8. Fiber optic Michelson interferometer.
The beam splitter with power split ratio e can be described by a two port scattering matrix with
the input-output relation written as [13]
(2.51)
By linearity, the response to an arbitrary input field can be assembled from the sum of the
responses to its Fourier components. Applying the scattering matrix to an input field E,(wo)ejt
44
and accounting for the reference and sample arm reflectivities RR ej(wOt) and Rs and propagation
lengths IR and is, one obtains the fields at the detector ER and Es as
ER(w,t)
=
j e(
e )E,(co)RR-j(2 iR'lRr)e-jO(w t)
AR ( )e-j(2 R-'R
-0e 1O)'t)(
-j(2,81S-(ot
)(
ois)- 22fl(2.52)t
(l- c)E,(o)Rse 2fs's-") - As(Wle(
Es (c, t) = j
where the front terms involving the coupler split ratio and input source power are grouped into
amplitude parameters AR and As. The time averaged photocurrent at the detector can be written
as [27]
'D
= hv
(2.53)
ER+Es 2)
217f
where q is the detector quantum efficiency, e is the electronic charge, hv is the photon energy
and 77 is the intrinsic impedance of the fiber core material. The detector response time is taken
to be much longer than the coherence time for a low-coherence source but much shorter than the
heterodyne signal oscillations. For a monochromatic source, the photocurrent evaluates to
iD
monochromatic
hv
[
71f
2
22AR
2
+ Re {EsE
}]
(2.54)
where
Re{E s E4 = ARAS cos(2GRlR ~28s s +#0(t))
(2.55)
The oscillating component of the interferences is seen to depend on the difference in phase
between the reference and sample arm light. In particular, for #(t) = 0 and equal propagation
constants in the reference and sample arm paths (p8R = s ), the beat term depends solely on the
path length difference between reference and sample arms. The product term ARAS in the
oscillating term provides the enhanced sensitivity of the heterodyne detection scheme. A very
small sample arm reflection As is amplified by the strength of the reference arm field AR . For
this reason, the power in the reference arm is often referred to as the heterodyne gain.
For a polychromatic, low-coherence light source, the oscillating component of the
heterodyne signal depends on the sum of the interference due to each monochromatic plane wave
and can be determined by integration of the cross-spectral term over the bandwidth of the light
source [27].
iD
o Re {ER(co, t)Es (o, t)*
) = Re
where
45
fS(co)e'd(o}
(2.56)
S(co) = As(co)AR (co)*
(2.57)
and
qp(co, t) = 28s (co)ls - 2,8R (c
R
+ #(w, t)
(2.58)
For the case where the coupler split ratio and the reference and sample reflectivities do not vary
across frequency, the term S(co) essentially represents the power spectrum of the source.
2.4.2
Coherence Gating
Consider the case when 8s =,8R = 8 and no additional phase shift
#(co, t)
is generated in the
reference arm. For a linear, non-dispersive medium the propagation constant /3 can be
represented by a first order Taylor series approximation about the center frequency co,
pi(co) = p8R (CO)
=
3(C)=
s
I(Co ) +
)'(Co)(0)
- 0o)
(2.59)
The phase difference yp(co, t) from (2.58) is determined by the path length mismatch Al = Is- iR
between reference and sample arms and can be written [27]
yp(co)= /p(co)(2A1)+p'(co,)(co - co )(2Al)
(2.60)
The expression (2.56) becomes
iD
o Re
edj'"Ar
f S(co - co0 )e
j(o-c")At" d(co-
coo)
(2.61)
2;T
where the A r is the phase delay mismatch and A r is the group delay mismatch defined as
.2A1 = 2A
A r, =
O
(2.62)
VP
and
Arg =P'(CoO) -(2Al) = 2A1
(2.63)
V9
The terms v, and vg are termed the phase and group velocities, respectively. They depend upon
the center frequency of the source spectrum and also the material properties of the medium, in
particular the index of refraction n . From (2.61) it becomes evident that phase delay term
46
creates a carrier frequency dependent on the center frequency of the source. As optical path
length varies, oscillations in the interference signal are generated. The heterodyne signal
envelope is the inverse Fourier transform of the function S(co) with respect to the group delay
parameter A z- * The group delay in turn is generated by path length variation between reference
and sample arms at a rate determined by the group velocity v,. Considering S(co) to be the
power spectrum, the heterodyne current then takes the form of an autocorrelation function in
accordance with the Wiener-Khintchin theorem.
Combining (2.52) and (2.57) with (2.61) and considering real sample and reference arm field
reflectivities constant in time and frequency, the heterodyne current can be written as
iD(Al) oc RRRs jF-' [S,(co)]l cos(coAr,)= RRRsG,
cos
(2.64)
where the source power spectrum S. (co) is related to the autocorrelation G, (A rg) by the
Fourier transform with respect to the group delay. The autocorrelation is a measure of the degree
of temporal coherence of the source. From the time-bandwidth principal familiar of Fourier
transform theory, it is clear that the width of the envelope of the interference signal decreases for
larger bandwidth, shorter coherence length light sources.
The envelope or axial point spread function of the interference signal is known as the
coherence gate. In optical coherence tomography (OCT), the width of this envelope sets the
axial resolution of the imaging system. To generate an image, the sample arm focus is scanned
laterally at a fixed depth in the tissue and the reference arm path length is varied at high speed.
The position of the reference arm effectively gates out light scattering from planes in the sample
to within the coherence length of the light source.
Many broadband sources of interest for low coherence interferometry can be approximated
by a Gaussian source spectrum. A normalized Gaussian power spectral density can be written as
[26]
S(co -
2
2
_)_=
(2.65)
where o is the standard deviation. When evaluated in (2.61) the Gaussian density generates a
Gaussian interference signal decribed as
iD c Re e 2r e
}
e2 ,
cos(cwAr,)
(2.66)
where the full width at half maximum (FWHM) of the interference signal in a free space
interferometer can be shown to be related to the center wavelength and spectral FWHM as
47
AlFWHM-
21n2
2
L
2
(2.67)
A2
This result is typically used for specification of the coherence gate or axial resolution of a low
coherence interferometry system. Note that the width of the coherence gate scales as the square
of the center wavelength, meaning longer wavelength sources require higher bandwidth to
achieve the same axial resolution. Figure 2.9 demonstrates the dependence of the coherence gate
on bandwidth and center wavelength for common near-infrared wavelengths used for imaging in
biological samples. Achieving a coherence gate of 5 um requires nearly 150 nm of bandwidth at
center wavelength 1300 nm but less than 60 nm bandwidth at a wavelength of 800nm.
25
---
----
800nm
1300nm
20-
-
?=15-
U-
10-
5-
0
0
100
50
150
200
Bandwidth (nm)
Figure 2.9. Coherence length versus bandwidth for NIR imaging wavelengths.
For a general non-Gaussian spectrum, the coherence gate can be evaluated by determining
the FWHM of its autocorrelation function. The time-bandwidth product of the Gaussian is
optimal, meaning for a given spectral width, non-Gaussian spectra will produce broader pointspread functions. Generation of sidelobes by non-Gaussian sources can also be a problem for
low coherence interferometry applications.
2.4.3
Effect of Group Velocity Dispersion
Group velocity dispersion (GVD) describes the phenomenon where different wavelengths of
light propagate through a material at different group velocities. For positive (normal) dispersion,
the index of refraction increases with decreasing wavelength, and for negative (anomalous)
dispersion, the opposite occurs. Most materials introduce positive dispersion. In low coherence
interferometry systems, the presence of a mismatch in the dispersive properties of the reference
and sample arm propagation paths leads to a reduction in the heterodyne signal amplitude, a
48
broadening of the axial point spread function, and a chirping of the carrier frequency.
GVD is
incorporated into the above analysis by expanding the propagation constant /s and R to
second order around the center frequency co. Including the reference arm phase term #(co,t)
the phase term y(co) from (2.58) then becomes
1
p(c)= #(co, t)+ pJ(co,)(2A)+ ,'(co)(co - wO)(2A) + -/p"(o)(o - co')2 (2AL)
(2.68)
2
where A/p"(co0 )= p, (co,,)- ,6(cqo) is the mismatch in the dispersion terms and AL is the length
of mismatch. The interferometric signal can now be written
iD c Re e-
AP
L
JS(Ct-cv)ed
"
e-)
e
~
-00
d(w
w
(2.69)
Note that correct choice of the reference arm phase shift can be used to compensate the quadratic
dispersion term P". In practice, the material dispersion in the reference arm and sample arm
must be carefully matched to ensure optimum resolution and sensitivity. For a Gaussian
spectrum, the interferometric signal becomes [27]
Arg2
iD oc
Re
IF(2L)
U
e
2F(2L)2
e -joAr
(2.70)
where o7 is the width of the Gaussian in a non-dispersive medium and F(2L)2 is a complex
function whose magnitude increases with increasing GVD mismatch. The presence of F(2L)2 in
the exponential leads to frequency chirp and envelope broadening of the interference signal. A
reduction in amplitude of the signal results from the inverse dependence of the scale factor in
front.
2.4.4
Detection Electronics
The detection electronics used for low coherence interferometry systems typically consist of
four principal stages: photodiodes, transimpedance amplifier, bandpass filter, demodulator.
Figure 2.10 shows a schematic. Dual balanced photodiodes convert the optical signals to
electronic currents and add them. The out of phase DC currents cancel each other and the in
phase oscillating components add. The transimpedance amplifier then serves as a current to
voltage converter. Because of their low-input current, field effect transistor (FET) input op amps
are typically used for monitoring photodiodes. After the transimpedance stage, the bandpass
filter isolates the frequency content of the signal from excess wideband noise and thus provides a
crucial component in improving the sensitivity of the system.
After the bandpass filter, several options are available for demodulation. The oscillating
interferometric output can be sampled directly with a high-speed data acquisition card and
processed by a DSP processor or by computer. A Hilbert transform technique is typically used
for envelope detection from a sampled signal. When the heterodyne frequency cannot be
49
sampled, however, analog demodulation must be used. Typically, an RMS converter rectifies
the signal and a low-pass filter removes the carrier. For high dynamic range signals, logarithmic
amplification is generally used as well. Analog demodulation can also be done with a quadrature
demodulator. A quadrature demodulator removes the carrier from the heterodyne signal by
mixing the signal with two sinusoids in quadrature. The baseband components for each mixed
signal are then low pass filtered and combined to provide amplitude and phase information about
the heterodyne signal.
Demodulator
Output
Interferometric
Output
RMS Converter
Bandpass Filter
Transimpedance
Amplifier
Lowpass Filter
Figure 2.10. Typical detection electronics used in low coherence imaging systems.
2.4.5
Noise Sources
The electronic noise spectrum after the transimpedance amplifier typically has three
dominant components: shot noise, thermal noise from the transimpedance resistance, and excess
intensity noise from the laser source. The double-sided noise power spectral density can be
written as [28]
Si
(Co)
=
S
(C),,,oS
+
S (C) e
+ S ( ),,,
=
e (i) + ey (i) 2 + 2kT(2.71)
where e is the electronic charge, k is Boltzmann's constant, R is the transimpedance resistance,
and y is a noise parameter which must be empirically determined. The square term (i) 2
represents the photocurrent power. Shot noise arises from current fluctuations due to the
quantization of light and charge and can generally be considered a white noise process with
mean (i). Thermal noise is generated by random motion of particles due to thermal energy in a
system and is associated with transfer of energy and temperature equilibrium between a resistor
and its surroundings. Excess intensity noise is a combination of all noise sources whose power
spectral density scales linearly with the mean photocurrent power. Examples include excess
photon noise and local oscillator noise [28, 29]. Use of dual balanced detectors can largely
eliminate the excess intensity noise by subtraction. Shot noise, however, cannot be eliminated
because the noise processes from the detectors are statistically uncorrelated and the variances
add when the photocurrents subtract.
The power in the noise process n(t) is given by the noise variance. For transimpedance
resistance R and noise equivalent bandwidth NEB the variance is written [26]
50
var {n(t)} = 2R 2 [Sin (co)]- NEB
(2.72)
The noise equivalent bandwidth is defined as the product of the low-pass and band-pass filters if
the bandpass filter was translated to the origin.
2.4.6
System Sensitivity
Low coherence interferometry systems are made to operate near the quantum sensitivity limit
by choosing system parameters so that the shot noise overwhelms all other noise sources.
Practically, this means combination of a very low noise analog receiver with sufficient reference
arm power to set the shot noise level above the receiver noise. The signal to noise ratio for the
system is defined as the ratio of the signal power Pn,,,,, to the power in the noise process n (t)
[28].
P.
SNR n1ise
S[R= 'ial =
P
signal
00
P.
signal
var tn(t) }.3
dc
2.73)
The shot noise from the reference and sample arm powers is determined from the DC
components IAR 2 and IAs 2 in (2.54). When near the sensitivity limit, power from the sample is
negligible compared to reference arm power. Using the definition for
DC photodetector current from the reference arm becomes
R
AR
provided in (2.52), the
(2.74)
1 6(1 - 8)E R 2
hv 2qf
R
where E is the input field amplitude, c is the coupler split ratio, and RR is the reference arm
reflectivity. Combining this with (2.72), the noise variance in the shot noise limit is written
nie
noise
The heterodyne signal power
hv
pignaI
1
277f
(
R
c(I-c) E
2-EB
(2.75)
ReR2 .NB
is determined from the time average of the oscillating
interference term of (2.54), labeled ibea, here for bookkeeping.
-2
'signal
(eat)b
R
=
7
1
2
hv 2qf
51
E2
RsRR R R 2
2
(2.76)
Dividing (2.76) by (2.75) now provides the shot-noise limited signal to noise ratio for a single
detector configuration.
SNRS
SNRsi
detector
PS
1 e(1-e)E R~ =__
16lcE2
- 17
hv 2- NEB
2 -NEB
hv 277f
i
=---
(2.77)
The signal to noise ratio depends only on the noise equivalent bandwidth and on the optical
power returning from the sample arm Ps. From the expression, it is clear that choice of a
coupler with even split ration e = 0.5 is optimal because it maximizes the amount of power
returning from the sample to the detector.
2.4.7
Dual Balanced Detection
Dual balanced detection improves system sensitivity by eliminating excess intensity noise
from the photodetector signal. It can also be shown that the use of dual balanced detection can
enhance the sensitivity in the shot noise limit. Consider the schematic in figure 2.11. The input
power is now normalized to provide the same sample power as the single coupler, single detector
configuration.
Source
EO/Va
Reference
G
D1
Sample
D2
Figure 2.11.
Dual-balanced interferometer configuration.
By extending the scattering matrix analysis from previous sections, it can be shown that the
oscillating components of the heterodyne interference signals at DI and D2 are in phase while
the DC components are out of phase. Addition of the two photocurrents at the transimpedance
amplifier results in cancellation of the DC currents carrying excess intensity noise and addition
of the oscillating heterodyne components. Attenuation is necessary at the high detector to ensure
that the DC powers match. It can be shown that the attenuation factor necessary is given by
(1- a)(1-)
Using this factor, the signal to noise ratio becomes
52
(2.78)
-
SNR dual
balanced
7
hv
2
1
7f
-
E2R 2
s
4. NEB
(2.79)
Comparing (2.79) with (2.77) and taking c = 0.5 for optimal performance, the enhancement of
the dual balanced configuration can be quantified.
SNRdual
balanced
=
2(1-a)
(2.80)
SNRsingle
detector
Compared to the single detector configuration, use of two even splitters e = a = 0.5 provides
equivalent performance and using a <0.5 offers enhanced performance in the shot noise limit
for fixed sample power. Note that for a <0.5, however, increased source power must be
supplied to obtain fixed sample power. The best case scenario is use of an optical circulator
instead of the first coupler. It provides complete transmission in both forward and reverse
directions and offers an SNR enhancement of a factor of 2 (3 dB).
2.5
Combined Confocal and Coherence Gating
The previous analysis for a low coherence interferometer included no provision for sample
arm focusing effects. From the derivations of section 2.3, it is clear that the presence of a
confocal microscope in the sample arm can have an important effect on the heterodyne signal at
high numerical aperture. Kempe and Rudolph have provided detailed analysis of confocal
microscopy and heterodyne microscopy for broadband laser sources [30, 31]. This section
considers a simpler approach that ties together previous analysis of confocal microscopy and
low-coherence interferometry from sections 2.3 and 2.4 followed by discussion of the
implications of combined confocal and coherence gating.
2.5.1
Heterodyne Signal for Combined Gating
Recall from (2.31) that the field at the detector in a confocal microscope
uconfocal (x,
y, z) is
given by the convolution of the field reflectivity function rs (x, y, z) with the confocal impulse
response [h, (x, y, z)] 2 . The confocal field response then interferes with the reference arm field
to give the heterodyne current. For an axially distributed reflectivity as would be found in
scattering media, the heterodyne signal can be written as an integral over the sample arm path
length ls in a similar manner to (2.32). Replacing Rs in (2.64) with the confocal response
modified reflectivity from (2.31) and integrating over the sample path the heterodyne current
becomes
53
iD
R)
G
dls RR
C
Lo
-0
Cos
-
I)G
V
Aconfocal/)}
V9
(2.81)
2cooAl
2A1
(,)(92h
lS)0[h,(ls)]2Go---jcos2W l)
=JdlsRRrSR~r
where Al
=Is
-
'R
is the difference in path between reference and sample arms and the (x,y)
dependence of u,onjoca, rs, and h, have been ignored for simplicity. The heterodyne component
is the convolution of the sample arm confocal response with the carrier dependent source
autocorrelation function. For a single scatter at location Is, with reflectivity Rs(lso), the
heterodyne current can be written as
iD (R)
oc RRRs(lso) Ic(s 0 )G0
2(I -1R )
2coo(IS -l I)
j COS))j
so
(2.82)
where, Rs is the sample reflectivity, and Ic(u) has been chosen to represent the confocal
intensity response for consistency with the results of section 2.2. By definition, IS= 0
For a point scatterer described by
corresponds to the position of the focus.
rs(Iso) = Rs(1,s)Sls -Is,) the confocal response simply reduces to the field impulse response of
the confocal system at position iso, [h, (ls)]2, which is given by the square root of (2.40). The
heterodyne current is now written as
in"t
RR
R
D
R
s
sin(uO / 4) G
uo / 4 _
SO(1Rj
0)
Kc
2(lso~IR)
V9
oR 2oso UR
2RiRs((Is
CS
V
where uO =8A IsOsin 2 (a/2) and sin(a) is the objective lens numerical aperture.
(2.83)
The
response for the fiber confocal microscope is given by (2.48).
2.5.2
Depth of Field and Transverse Resolution
From (2.82) it is evident that control of the position of the confocal and coherence gates is
decoupled. The confocal gate is set by the focus position, in this case iso. For fixed focus
position, however, the coherence gate is set by the reference arm path length 1R. Maximum
heterodyne signal current occurs when the reference and sample arm path lengths match 'SO = R
such that the coherence and confocal gates precisely overlap. When a mismatch between the
gate positions exists, the action of the confocal gate on the coherence gate can dramatically
reduce the heterodyne signal current. The sensitivity of heterodyne signal to the relative position
of the gates scales with the numerical aperture of the objective lens. Figure 2.12a demonstrates
this principle. For high numerical aperture, the confocal parameter b is quite small and results
in a small depth of field for imaging. When the difference between the reference and sample arm
54
path lengths increases beyond the confocal parameter, the confocal and coherence gates no
longer overlap and the heterodyne signal amplitude is diminished significantly.
Low-NA
High-NA
'Alc
b
l
b
A=Is-IR
Figure 2.12. Effects of focusing on depth of field.
Because of this limitation, optical coherence tomograpy (OCT) systems operate with
relatively relaxed numerical aperture focusing, shown for comparison in figure 2.12b. In OCT,
the focus position is set to a fixed depth in the sample and the reference path length is scanned.
When the confocal parameter is sufficiently large, the backcoupled intensity from the confocal
microscope is nearly constant over the range of the depth scan and an adequate depth of field is
preserved in the images. The cost of the relaxed confocal gate, however, is larger focal spot size.
Hence there exists a fundamental tradeoff in OCT between depth of field and transverse
resolution.
2.5.3
Optical Coherence Microscopy for High Resolution Imaging
Optical coherence microscopy can overcome the tradeoff between depth of field and
transverse resolution by using an alternate scanning method to conventional OCT. OCM
techniques fix the reference path to match the sample path, allowing use of high numerical
aperture objective lenses to provide high transverse resolution. An enface image is then mapped
out by raster scanning the beam on the sample as is done in confocal microscopy.
OCM systems do not require group delay scanning as in OCT. In order to generate a
heterodyne interference signal, however, reference arm phase delay scanning is still necessary.
Without it, it is clear from (2.64) that when the reference and sample paths match the oscillating
heterodyne signal current disappears. Phase delay scanning in the reference arm can be
accounted for by considering again the wavelength dependent phase shift term #(co,t) in (2.52).
Expanding #(co, t) about the source center frequency in a first order Taylor series approximation
gives
55
#(Co,
(2.84)
(
t) = 0(o, t) + 0'(Co, t)(o - co)
Incorporating (2.84) with (2.60) the heterodyne signal current is now evaluated via (2.56) to be
'D
where
OeRe
#(coo , t)
-
e-A""')'e-i,(">2A G
o,,xo,-(O)e-jpgox(-m)A'
represents a phase delay term and
#'(coo,
dW(2o
(2.85)
t) represents a group delay term. For
pure phase modulation, the group velocity term #'(wo, t) should be zero. Considering again real
sample and reference arm reflectivities constant in time and frequency, the heterodyne current
from (2.64) becomes
iDSo
RsRRGo
cos(2p(wo)Al+#(oot))
(2.86)
In OCM, the reference arm path length is set to match the focus position defined as Is = 0. For
an axially distributed sample reflectivity characteristic of a scattering medium, the OCM
heterodyne signal is given by the integration of (2.86) over all sample paths with 1R = 0.
Incorporating the confocal response, the signal takes the form
iD (t) OC
dlsrs (s )
[h,(lS )]2]GO
2ls cos [2/3(coo)ls + (co ,t)]
(2.87)
With reference arm fixed, the integral is no longer a convolution of the confocal response with
the autocorrelation function. Note from the term 2pl(cqo)ls represents the pathlength dependent
phase shift in the interference component of each scatterer. The summation over the distributed
scatterers results in a reduction of the heterodyne amplitude because the interferences do not add
coherently.
For a single scatterer located at position ISO the heterodyne signal simplifies to
'D
(t)
oc RRRs (ls)
so cos[23(wo
0 )lsO+$(o 0,t)]
Ic(ls)G
(2.88)
When the sample is in focus, the coherence and confocal gates overlap and the amplitude of
the oscillating interference signal is maximum. Note that as the position of the sample iso
changes with respect to the focus ls =0, the amplitude is subject to the multiplication of the
coherence and confocal gate point spread functions.
The heterodyne signal varies in time with the reference arm phase delay scan #(co, t).
Raster scanning the beam in an en face pattern modulates this heterodyne signal by the
56
reflectivity profile of the sample. A plot of this signal amplitude versus beam position makes up
an OCM image.
To obtain a pure sinusoidal oscillation, a phase ramp linear in time must be imparted to the
reference arm field. Section 2.6 will discuss the use of a grating phase delay line for this
purpose.
2.5.4
Path Length Scaling with Focal Position in Tissue
From (2.58) it is clear that the phase difference between reference and sample fields that
determines the heterodyne signal depends on the optical path length (OPL), which is defined as
OPL = n(w) -l
(2.89)
where 1 is the physical (free space) path length and n (co) is the refractive index of the medium
of interest given by the relation 8 (co) =n (co)
-.
.
For simplicity, the frequency dependence of
C
n is ignored here. Results derived thus far have assumed that propagation constants in the
reference and sample arms match, 8 = /s =6R, meaning the optical path length is equivalent to
the free space path length. When the reference and sample refractive indices match, ns = nR , the
reference arm OPL remains equal to the sample arm OPL as the focus is translated deeper into
the sample. This does not hold true when the refractive index of the sample is different from that
of the reference arm. In this case, a physical thickness of amount S over which ns # nR
produces an OPL mismatch of
ALOPL
(nsnR
(2.90)
)-An5
Accounting for the mismatch and ignoring the frequency dependence of the index of refraction,
the phase difference between reference and sample arms is
CO
C
C
C
yp(w,t)=2 -nAl+2
(2.91)
(An)Sl+#(w,t)
where the first term represents a physical mismatch between reference and sample arm paths of
common index no without the index effects of the sample and the second term accounts for the
effects of a change in index over S. With this phase function, the heterodyne signal for OCM
from (2.88) becomes
iD(t) oc RRRs (iso ) Ic(s)G
2nils +2Anti
c
C
cos[2(cq)sO + 2 cooAns +#(coo,t)] (2.92)
c
OCM is critically sensitive to overlapping the confocal and coherence gates by matching of
the reference and sample arm optical path lengths. Any OPL change due to focusing into a
sample with a mismatched index leads to a reduction in heterodyne amplitude. To optimize
57
heterodyne signal, the reference arm path must therefore be coordinated to the focus for each
focal depth into tissue.
2.5.5
Enhanced Gating Effects in Scattering Media
The relative merits of combined coherence and confocal gating with heterodyne detection
were mentioned with respect to previously published work in chapter 1. Key points are
reiterated and developed here for clarity. The heterodyne detection process provides nearly shotnoise limited detection using optical amplification via a reference arm field [32]. This feature
makes heterodyned detection more sensitive to weakly scattering objects than direct detection
alone. Moreover, correlation with a reference arm signal allows a mechanism for correcting
chromatic and spherical aberration. Phase modification of the reference field can be used to
cancel aberration produced in the sample arm optics or in the sample itself. Use of identical
sample and reference arm optics, for example, eliminates aberrations from the optical system
[33].
The combined effect of the coherence gate and confocal gate on axial dependence of the
heterodyne signal amplitude provides enhanced rejection of unwanted scattered light in an OCM
system compared to a confocal system alone. Together, combination coherence and confocal
gating can provide improved depth discrimination and image contrast [31, 34, 35]. The
improvement is evident first in the functional form of the axial point spread functions. The
coherence gate for a typical low coherence source approaches a Gaussian in form, which falls off
with axial position much faster than the typical sinc 2 decay of the confocal response. In highly
scattering media, the Gaussian function provides better rejection of out of focus scattered light to
overcome the exponential decay of incident light.
Axial sectioning is also enhanced by distinct mechanisms of rejection for coherence versus
confocal gating. Confocal gating depends on the position of the scatterer relative to the focus
and is therefore purely a spatial gating technique. The coherence gate, however, selects for path
length of returning photons. The value of using both rejection mechanisms can be understood by
returning to the schematic representation of heterogeneously scattering medium shown in figure
2.2. The single backscattered light contains the image information. Multiply scattered light
from planes outside of the focus degrades image contrast and ability for depth discrimination.
Confocal microscopy provides no intrinsic mechanism for removal of multiply scattered light.
Light that scatterers into the focal volume and then into the collection angle of the objective can
fall within the confocal gate and obscure the single scattered component. Coherence gating,
however, provides path length discrimination against the multiple scattered component and
thereby enhances confocal gating. Similarly, the presence of the confocal gate helps improve
coherence gated sectioning. The multiply scattered light that happens to traverse the correct path
length and fall under the coherence gate will be rejected by the confocal sectioning.
Izatt et. al. placed some quantitative limits on the enhancement provided by combined
coherence and confocal gating in scattering media using single-backscatter theory [34]. The
confocal signal is considered to be the integral of the mean heterodyne signal, or equivalently,
the sample arm DC photodetector current as expressed in (2.32). Confocal microscopy fails
when the sum of light intensity from all planes other than the focus dominates the level at the
focus. In terms of scattering mean free paths (MFP) this depth limit zconjocaI is
58
2 2
z confocal :<,2- n [r
D (NA)2
_ 4n2A2
L0
(2.93)
1
where NA is the numerical aperture of the objective, no is the index of the scattering medium,
D is the sample arm beam diameter on the objective, and M is the magnification due to the
lens. The limit for heterodyne detection is the quantum shot noise sensitivity limit, expressed as
z
Zcoherence
! 1 InE
2
[2hv
rpbLc
NA2
2L (NA)
no
I
(.4
(2.94)
where pb is the backscattering coefficient and Lc is the coherence length. OCM improves upon
confocal microscopy between these depth limits
Zconfocal
-
z
:
Zcoherence
(2.95)
Based upon these calculations, confocal microscopy provides imaging to depths of 5-9 MFP
while OCM extends the range to between 15-20 MFP. Combined coherence and confocal gating
can thus improve image depth by 2-3 times over standard confocal microscopy.
The results from the single scattering analysis point out that the confocal signal can be
largely dominated by high scattering at or near the surface. In this case, use of even a moderate
coherence gate for imaging deeply in a scattering medium can eliminate this surface component
and improve imaging depth range.
The single scattering model has been shown to agree with experimental results for a
homogeneous scattering phantom [34]. When imaging deeply in heterogeneously scattering
media such as biological tissue, multiple scattering effects become important. Multiple
scattering causes the beam to spread in scattering media and can significantly degrade the
transverse coherence of the beam. Resolution and maximum probing depth are thereby
decreased. Consideration of multiple scattering in OCT has been achieved using the extended
Huygens-Fresnel principle and mutual coherence functions [36-38].
Schmitt considers a
maximum probing depth zmax in OCT in the presence of multiple scattering as the depth at which
the strength of the multiple scattered component of collected light equals the strength of the
single scattered component [36].
Zrax
In K1+ 2R2
~
p2
2p,
(2.96)
In this expression, R represents the beam radius on the sample and p is a measure of the
transverse spatial coherence of the beam.
59
2.5.6
Operating Regimes for OCM
As with the position of the gates, the width of the confocal and coherence gates can be
independently controlled. The duration of the confocal gate is set by the focusing characteristics
of the sample arm optics, described by the numerical aperture (NA). Higher numerical aperture
leads to a shorter confocal gate and vice versa. The width of the coherence gate on the other
hand is set by the bandwidth of the light source, independent of the sample arm optics. Larger
bandwidth produces a smaller coherence gate. Considering this independent control, four
distinct operating regimes for high resolution optical coherence microscopy can be defined based
on the relative sizes of the confocal and coherence gates.
V
-
X-ct:
Short Confocal
Long Confocal
Short Coherence
Short Coherence
Short Confocal
Long Confocal
Long Coherence
Long Coherence
aac
Figure 2.13. Operating regimes for optical coherence microscopy.
Short and long are defined here relative to accepted axial sectioning requirements for
imaging in turbid biological tissues with confocal microscopy. Deep imaging of singly
backscattered light in highly scattering tissues requires an axial section thickness significantly
smaller than the mean free path of scattered photons in the specimen, which is typically 20 - 100
um in tissues of interest. Empirical results with in vivo laser scanning confocal microscopy have
defined a suitable section thickness for achieving high-contrast confocal images of cellular
structure to be less than 5 um [16]. This is less than the typical thickness of a single layer of
cells and consistent with standard section thickness used in histologic analysis.
For deep tissue imaging with OCM, the section thickness can be set with the coherence gate,
the confocal gate, or a combination of both. Plotting equation (2.42) for the confocal gate
against the coherence gate defined in equation (2.66) illustrates the interaction of the gates. The
extreme limits for OCM are compared in Figure 2.14. In 2.14a, a 3um coherence gate is plotted
with a 3um confocal gate. This configuration represents the high-resolution limit of OCM. It
clearly provides adequate axial and transverse resolution for cellular imaging, and should in
theory provide very high contrast. This configuration, however, is practically very difficult to
60
implement due to the considerations discussed in sections 2.5.2 and 2.5.4. Even a very small
displacement of the coherence gate with respect to the confocal gate generates a dramatic
reduction in heterodyne signal. In light of the turbid refractive index in biological tissue, overlap
of the gates will be difficult to maintain.
Figure 2.14b demonstrates the opposite limit where the confocal and coherence gates are
relatively long, in this case 30 um each. This operating regime clearly does not provide adequate
sectioning ability for high-contrast imaging in highly scattering media. On the positive side, it
does not suffer from the extreme sensitivity to displacement of the confocal and coherence gates.
Path length changes of tens of microns are needed before the heterodyne signal is lost.
Nonetheless, the lack of adequate sectioning prevents practical implementation of this
configuration for cellular imaging applications.
Short Confocal / Short Coherence
Long Confocal
1.2
1
1
0.8
0.8
0.6
0.6
0.4
0.4
0.2
0.2
-50
0
50
/
-50
-50
0
0
50
Axial Position (urn)
Axial Position (urn)
-50
50
0
0
50
0-
-20
-0
/ Long Coherence
1.2
-20
-40
IM
0o
.2
-60
-80
-40-60r
-80F
0%^^
-100'
B
A
Figure 2.14. Extreme operating regimes for optical coherence microscopy. The coherence
gate is represented by the dashed line and the confocal gate by the solid line.
The compromise regimes for OCM imaging are shown in Figure 2.15a and 2.15b. The short
confocal gate / long coherence gate regime displayed in 2.15a can essentially be thought of as
coherence-gated confocal microscopy. The confocal gate dominates in setting the section
thickness while the coherence gate serves to enhance the rejection of out of plane scattered light.
Previous work on optical coherence microscopy has been performed in this regime. Recall from
the discussion of section 2.5.5 that even a moderate coherence gate can knock out the surface
scattering that degrades images of confocal microscopy. Coherence-gated confocal microscopy
has been demonstrated for deep tissue imaging to depths of 600 um, well beyond that of standard
61
laser scanning confocal microscopy [39]. The sensitivity of the heterodyne signal to path length
mismatch is reduced in this configuration. The small confocal gate can essentially shift around
under the larger coherence gate without losing significant heterodyne amplitude. Note that as in
confocal microscopy the dominant sectioning power is determined in this case by use of high
numerical aperture objective lenses. Such lenses are generally bulky and contain multiple
elements, making development of miniaturized probes for confocal microscopy and coherence
gated confocal microscopy difficult.
Long Confocal / Short Coherence
Short Confocal / Long Coherence
1.2
1. 2
I
0. 8-
0. 8-
E' 0. 6-
\
0. 6 -/
0. 4 -/
0. 4 -
0. 2 -
0.2 -
0 -----
L
-50
0
0
50
-50
Axial Position (um)
-50
0
0
50
Axial Position (urn)
-50
50
0
0
-20-
-20
-40
-40
2
0 -60
0
50
-o
-60
-80
-80F
-inn[
A
B
Compromise operating regimes for optical coherence microscopy. The
Figure 2.15.
coherence gate is represented by the dashed line and the confocal gate by the solid line.
The final operating regime shown in Figure 2.15b uses a short coherence gate in conjunction
with a relatively long confocal gate. This configuration uses the coherence gate to set the axial
section thickness and might suitably be called high-resolution optical coherence microscopy to
distinguish it from confocal microscopy. The sample arm optics in this configuration provide a
high transverse resolution but only a relatively weak sectioning power. Again, the relatively
longer confocal gate reduces the sensitivity of the heterodyne signal to pathlength mismatch.
The coherence gate can now effectively shift within the window provided by the confocal gate
by tens of microns before the heterodyne signal is lost. Importantly, this configuration reduces
the numerical aperture requirement for achieving cellular imaging. Recall the comparison of
axial and transverse resolution in figure 2.5. Note that the transverse resolution of the confocal
microscope remains a few microns even for relatively small NA values. It is the rapid loss of
axial resolution with lower NA that destroys confocal images. High-resolution OCM can take
62
advantage of this result to achieve cellular resolution at relaxed NA, thereby reducing the design
constraints for miniaturized probes needed in clinical applications. In addition to providing the
enhanced gating effects discussed in section 2.4.5, this operating regime can enable cellular
imaging in situations when confocal microscopy cannot.
This short coherence gate / long confocal gate regime has not yet been demonstrated in
published literature for in vivo imaging of human tissues. Implementation requires broadband
light sources and a broadband OCM system. Chapters 3 and 4 discuss the development and
demonstration of a broadband OCM system design suitable for cellular level imaging in the short
coherence gate / long confocal gate regime.
The plots in figures 2.14 and 2.15 are shown in both log and linear form. The log plots point
out particularly well the enhanced sectioning effects provided by the functional form of the
coherence gate compared to the confocal gate. For intuition, the signal in confocal microscopy
and in coherence-gated imaging can be considered to be the area under the gate in the plot. The
Gaussian envelope falls off dramatically faster than the sinc 2 dependence of the confocal gate,
thereby improving rejection of out of plane scattered light.
2.6
Phase Delay Line Modulator
The heterodyne signal in OCM depends on the introduction of a phase delay term #S(co,t) to
the reference arm field. A phase delay line similar to those used in optical coherence
tomography systems is used in this thesis to provide phase modulation. Figure 2.16 illustrates
the geometry for analysis.
incident
galvo
f\lens
mirror
grating
xo
-..
I
I
x(A)
'..
y(t)
m =u
specular
I
0. 1
L
Figure 2.16.
OCM.
I
4f
Schematic of grating phase delay line used to provide phase modulation for
63
Tearney et. al. initially demonstrated the delay line for high speed group and phase delay
scanning [40]. The grating and lens combination effectively act as a Fourier transformer. A
scanning galvanometer mirror then imparts a wavelength dependent phase shift to the dispersed
spectrum. This phase shift is mapped to a group and phase delay when the beam is retroreflected
and transformed back by the grating and lens. By choosing appropriate parameters for the
grating, lens, and position offset of the spectrum on the scanning mirror, it was subsequently
shown by Zvyagin and Sampson that the group delay could be set to zero and pure phase
modulation achieved [41]. This section explains the phase and group delay characteristics of the
modulator.
2.6.1
Grating Conventions and Notation
The equation that determines the diffraction characteristics of a grating is given as
m(A))
d
sin(0,)+sin(
(2.97)
where 0, is the incident angle, O(A) is the diffracted angle, m is the diffraction order, and d is
the ruling spacing of the grating. By convention, angles measured counterclockwise from the
grating normal are positive and clockwise angles are negative. Order numbers m that are
measured counterclockwise from the m = 0 specular reflection are considered positive. Gratings
are typically blazed to operate with maximum efficiency in the first order in the autocollimating
or Littrow configuration, shown in figure 2.17a. The delay line configuration, however, requires
that the grating be used with a reversed blaze direction in order to pass through the lens to the
scanning mirror. This is shown in figure 2.17b. The grating blaze determines only the efficiency
of the light throughput in a particular diffraction order, not the dispersion characteristics or the
spacing of the orders. The Littrow angle is given by Oi = 0 B where OB is the blaze angle of the
grating. From the grating equation above, the Littrow incident angle can be written as
0B = sin-1
2d
2d)
(2.98)
With the reversed grating, equivalent efficiency is achieved when the first order diffracts
orthogonal to the blaze face. This incident angle is given by the grating equation to be
Oi = sin-
(3,
2d
(2.99)
The group and phase delay characteristics for the device are identical whether autocollimating or
reversed configuration is used. To allow for positive angle notations that simplify equations,
operation near Littrow angle will be subsequently assumed.
64
A) Autocollimating
B) Phase Delay Line
Figure 2.17. Autocollimating and reversed grating configurations.
2.6.2
Phase and Group Delay Equations
Assuming L = f , separation between collimated wavelength components A and A, after the
lens can be written
x(L)=
f tan(0(A.)-0 0 )
(2.100)
where 00 is the diffracted angle of the source center wavelength 2k.
phase shift produced by the tilted mirror is
(1, t)= -2,8z(A)
The wavelength dependent
(2.101)
With the spectrum offset from the center of the scanning galvanometer mirror by x0 the
wavelength dependent pathlength is written as
z(A)
=
[x +f tan(0(A)-0
0 ) ]tan(y(t))
(2.102)
where xO > 0 corresponds to the mirror side where z (A) increases as A increases.
wavelength dependent phase shift is now described as
# (A, t) = -28f tan (y(t)) tan (O()
Recasting as a function of frequency, it becomes
65
-00)
The
(2.103)
2cof tan (
(O))tan
(-
+
(2.104)
Using the grating equation (2.97) and evaluating the derivative for arbitrary incident angle, the
phase and group delay equations can be shown for free space to be
O
_q(cw,t)
2x, tan(y(t))
' ---
C
C~OO
ao(co, t)
2x0 tan(y(t))
1Ooocd
2fA, tan(y(t))
I-
(2.105)
AO-sinO,
d
These expressions are typically reduced further using small angle approximations tan(x)~ x,
sin (x) x, and cos(x) ~1 together with the assumption of Littrow angle given by (2.98). In
addition, the delay lines are generally used in double-pass configuration both to double the delay
characteristics and to reduce backcoupling modulation resulting from beam walkoff that occurs
when the mirror scans. The reduced equations for double-pass configuration then become
4y (t) xO
C
Tg
4y (t )xO
(2.106)
4y (t) f ,
cd
4y(t)
c
(216
From the equations, it is seen that the scanning mirror generates phase and group delay. For
a linear velocity scan, a linear phase delay results. For position xO =0, however, there is no
offset of the beam on the galvo mirror and the phase delay is zero. Furthermore, for correct
choice of grating and lens parameters d and f , there exists a position x0 where the group delay
is zero. This is the operating point used in OCM for phase modulation. Figure 2.18 illustrates
the delay parameters (2.106) as a function of mirror offset. For a center wavelength of 800 nm,
grating groove density 1 / d of 80 lines/mm, mirror focal length f of 5 cm, and scan angle of 3.6
degrees, the zero group delay offset is located at 3.2 mm. Note that relatively low dispersion
gratings must be chosen to achieve an offset that is small enough to fit on available galvanometer
mirrors, which typically have mirror width of around 1 cm for high-speed versions.
66
Group Delay vs. Mirror Offset
Phase Delay vs. Mirror Offset
1.4
1.2
2
E
E
E
1.5
C
0.8
0.6
1
0.4
0.5
0.2
.5
U.
.5
0
mirror offset (mm)
5
0
mirror offset (mm)
Phase/Group Delay Ratio vs. Mirror Offset
Electronic Center Frequency and Bandwidth
1snn
15
-fo
---BW
1400
10 -
1200 .
5-
-fB-
1000 -
0
800Cr
4) 600
4h
400
0
-5-10
200-
-1-5
-5
___
.5
5
0
mirror offset (mm)
0
5
mirror offset (mm)
Figure 2.18. Phase delay line characteristics for X = 800 nm, f = 5 cm, l/d = 80 lpmam, y =
3.6 degrees. and x.= 3.2 mm.
Also plotted in figure 2.18 are the electronic frequency and bandwidth that result from delay line
operation with the given parameters. At the zero group delay offset position, the center
frequency is 1 MHz but the bandwidth reduces to zero indicating a pure sinusoidal heterodyne
signal. The analytical expression for electronic Doppler frequency fD comes from the equality
in (2.62) which can be written as
fD =
v~/3(>o
0)
2;r
-
(2.107)
1"
where v, is the time derivative of the free space phase delay path difference Al, = cr
aAl(t)
at
(2.108)
Using expression (2.106) for r, , the Doppler frequency is given as
4x, ay(t)
A,
67
at
(2.109)
The bandwidth in turn relates to the free space group path difference Alg = crg by the relation
Af =
(2.110)
0
where the group velocity is defined as
V = OAlg (t)
(2.111)
at
Using expression (2.106) for rg , the bandwidth becomes
Af=A2
z2x
f
d
(
1
(2.112)
a7(t)
at
From figure 2.18 it is also clear that an asymmetry exists in the group delay characteristics.
The asymmetry results from the fact that the wavelength dependent phase shift increases with
increasing wavelength on one side of mirror and decreases on the other. The phase delay, by
contrast, is symmetric with respect to the zero offset position. Figure 2.19 illustrates the
simulated heterodyne signal at symmetric mirror offset operating points for the above delay line
parameters. A bandwidth of 200 nm is assumed with a linear mirror scan.
Xo =
Xo = -3.2 mm
1.
-0.5
Xo =3.2 mm
-
-0.5-
-1
0.085
0 mm
0-
1
-0.5
-11 1'
0.09
0.095
0.1
0.105
time (mS)
0.11
0.115
0.085
-0.09
0.095
0.1
0.105
time (mS)
-1
0.11
0.115
0.085
0.09
0.095
0.1
0.105
time (mS)
0.11
Figure 2.19. Distinct operating points for grating phase delay line. Simulation parameters
are k = 800 nm, f= 5 cm, l/d = 80 lpmm, y = 3.6 degrees.
68
0.115
2.6.3
Dispersion Compensation
By varying the spacing L between the grating and lens, the delay line can be used for
dispersion compensation. The second order phase term can be described as a function of
deviation of distance L from the focal length f as [10]
[82(ot)]
aC2
/ (L - f)
3
3
2d2
[cos(0 0 )]
=- C
(2.113)
Interestingly, it can also be shown by differentiating (2.104) that for L = f the second order
phase varies with time, implying non-uniform dispersion characteristics of the heterodyne signal
over the duration of the modulator scan.
D2g(C,t)
w
L
ftan(y(t))
_3
Lcos (O)]
= C 2d2
10
2
iL_
69
(2.114)
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38.
39.
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Advances in Instrumentation and Comparison With Histology. The Journal of
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Gu, M., C. Sheppard, and X. Gan, Imageformation in afiber-opticalconfocal scanning
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Wang, H.-W., J. Izatt, and M. Kulkarni, Optical Coherence Microscopy, in Handbook of
Optical Coherence Tomography, B. Bouma and G. Teamey, Editors. 2002, Marcel
Dekker: New York. p. 275-298.
Hee, M.H., Optical coherence tomography of the Eye, in ElectricalEngineeringand
Computer Science. 1997, Massachusetts Institute of Technology: Cambridge, MA. p.
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71
72
Chapter 3
OCM System Development and Characterization
3.1
Overview
To explore the advantages of OCM imaging with a short coherence gate, a system that can
support large optical bandwidth is required. This chapter describes the design and development
of broadband OCM systems for real-time, in vivo imaging at wavelengths of 800 nm and 1300
nm. Figure 2.1 shows the schematic system diagram for reference. Section 3.2 discusses
important design goals for in vivo OCM imaging systems. Broadband light sources used for
generation of short coherence gates are described in section 3.3 and discussions of the backbone
interferometer, reference arm phase modulator, and sample arm optics follow in sections 3.4, 3.5,
and 3.6, respectively. The receiver specifications and data acquisition scheme used for highspeed imaging are described in sections 3.7 and 3.8. Finally, characterization of system
sensitivity and axial resolution are described in sections 3.9 and 3.10.
3.2
Requirements for In Vivo Cellular Imaging
OCM systems for real time, in vivo cellular imaging in human tissues face several challenges.
Some of these considerations are the same as those constraining design of in vivo laser scanning
confocal microscopes [1, 2]. Axial section thickness must be sufficiently small to resolve single
scattered light deep in tissue. Adequate lateral resolution to visualize cells and their nuclei must
be achieved. Furthermore, a sufficient field of view should be maintained so that aspects of
cellular arrangement and tissue architecture can be surveyed as well as individual cells. Other
requirements include high frame rate to visualize physiologic motion such as blood flow and to
combat against unwanted motion artifact from movement of the sample. For stability on the
micron size scale of cells, mechanical stabilization schemes are generally required. Finally, deep
tissue imaging requires use of near infrared (NIR) wavelengths for increased penetration and
long working distance objective lenses to accommodate the imaging depth.
Several unique constraints also apply to OCM systems for short coherence length imaging.
First, the systems require use of broadband, high-power laser sources. Use of broadband NIR
laser sources in turn requires a high-speed, achromatic phase modulator and sample arm optics
optimized for the NIR. In addition, high speed OCM requires a low noise radio frequency
analog receiver for detection and demodulation of the heterodyne signal and a path length control
mechanism to combat against sample induced delay mismatch.
Finally, for transition of imaging technology to clinical applications, design of compact probe
technology is required. Handheld and endoscopic devices enable access to tissues unavailable
for imaging with bulky microscope systems.
The following list summarizes initial performance goals for a high resolution OCM imaging
system.
73
"
"
"
*
*
"
"
"
"
"
Broadband sources and system components for use in NIR wavelength range 800 nm
- 1300 nm
Lateral resolution of at least 3 - 4 um
Axial section thickness smaller than 5 um
Minimum field of view of 100 um x 100 um
Imaging depth approaching 1 mm or more
Working distance at least 1 mm
Safe irradiance levels
Stable imaging in the presence of live tissue motion
Sample & reference arm path length coordination
Compact and miniaturized devices for clinical application
Subsequent sections will discuss implementation of technology for achievement of several of
these goals.
3.3
Broadband Light Sources
Broad bandwidth light sources in the near infrared wavelength regime have been under
development for telecommunications systems for many years. More recently, design of new
sources has been a major focus for research on optical coherence tomography [3]. Source design
criteria include wavelength, bandwidth, single transverse mode power, and stability. The
bandwidth directly determines the size of the axial coherence gate. To image with maximum
sensitivity, sources should provide enough power to operate near the exposure limit for tissue.
Exposure depends on the wavelength, imaging speed and the focusing conditions and is typically
in the range of 10-20 mW for in vivo systems. Dual-balanced low coherence imaging systems
have low throughput due to the fiber couplers and the loss in the sample arm optics. Given these
constraints, fiber coupled input power in the range of 20-50 mW or higher are necessary for
OCM systems depending on the precise interferometer and sample arm configuration.
Semiconductor superluminescent laser diode (SLD) sources have largely been the standard for
low coherence medical imaging applications to date. Early sources for OCT had center
wavelengths around 850nm with near 20 nm bandwidth, providing axial resolution of
approximately 15 um. These sources, however, were limited to below 1 mW output power, and
were not sufficient for real time imaging. A high-power superluminescent diode laser (SLD)
source at 1300 nm was developed by AFC Inc. in 1997 to produce over 60 nm bandwidth and 20
mW output power. This source has become the standard for clinical OCT systems. The need for
higher power, broader bandwidth sources has necessitated the use of femtosecond laser
technology. Short pulse lasers provide bandwidth inversely proportional to the pulse duration.
In the femtosecond regime, bandwidths of hundreds of nanometers can be achieved with 50 - 100
mW or more of output power. The remainder of this section discusses the laser sources used for
OCM imaging in this thesis as well as sources available for future OCM imaging.
3.3.1
Semiconductor Superluminescent Diode Laser Source at 1300 nm
Initial OCM imaging results were obtained with a SLD source providing ~ 65 nm bandwidth
at 1330 nm with 20 mW output power. Figure 3.1 demonstrates the source spectrum used for
imaging. The nearly Gaussian form of the spectrum provides a low-noise, echo free coherence
74
gate and the long 1330 nn wavelength allows for increased image penetration. The coherence
gate for 65 nm is approximately 12 um in free space, which is sufficient for coherence gated
confocal microscopy. In operating conditions where the coherence gate dominates axial
sectioning, however, a 12 um resolution is insufficient for cellular imaging deep in tissue.
1.2
1
0.8
0.6
F-
o0.4
I-
E
0.21-
1100
1150
1200
1250 1300 1350
wavelength (nm)
1400
1450
1500
Figure 3.1. Spectrum produced by amplified SLD source used for OCM imaging. Center
wavelength is 1330 nm with bandwidth 65 nm.
The SLD source provides reliable turnkey operation in a variety of environments, easing its
integration into a clinically useful system. Nonetheless, short coherence gate OCM imaging
requires femtosecond laser sources until further improvements in SLD technology emerge.
3.3.2
Modelocked Ti:A13 0 2 Femtosecond Laser Source at 800 nm
Investigation of short coherence length OCM imaging was performed with a modelocked
Ti:A130 2 solid state laser producing femtosecond pulses at 800 nm. The laser was originally
designed and demonstrated by Morgner et. al. for generation of sub-two-cycle optical pulses with
bandwidth in excess of 400 nm [4]. Pulses of ~ 5.4 fs duration were produced at 90 MHz with
an average power of 200 mW. Figure 3.2 illustrates the laser cavity design. The Ti:A130 2
crystal is pumped by a continuous wave (CW) argon-ion laser producing 5 W ouput power. The
resonator is a standard Z-fold design with the crystal oriented at Brewster's angle. Low
dispersion prisms and novel double chirped mirrors are used for broadband dispersion control.
CaF 2 prisms intracavity provide positive dispersion to reduce the amount of negative third-order
dispersion that must be balanced by the DCM's. Fused-quartz prisms extracavity provide further
dispersion compensation to reduce pulse width.
This laser source has been used extensively for ultrahigh resolution OCT studies [5-7]. The
spectral bandwidth for imaging can be controlled by adjusting intracavity prism insertion. The
adjustment maintains average output power. To further shape the spectrum to obtain the highest
axial resolution, an extracavity spectral shaping apparatus is included between the quartz prism
pair. Three individually controlled fibers provide spatial line filters for narrow bandwidth fine
tuning of the spectrum while a slit aperture can be used to shape the high and low ends of the
spectrum and to set the center wavelength. Extracavity shaping reduces the average power of the
75
output beam to the interferometer and must therefore be done considering the power
requirements of the particular imaging application.
Interferometer
OC
Mo
,I
M4
M,
3
MP
X M2
PI
L
P2
RB
FS
P4
M6
Figure 3.2. Cavity setup for modelocked Ti:A120 3 femtosecond laser. Pump lens L, crystal
X, curved mirrors M2 and M3, flat mirrors MO, MI, and M4-M7, output-coupling mirror
OC, intracavity prisms P1 and P2, and extracavity prisms P3 and P4, extracavity spectral
shaping fibers FS and razor blade RB.
1.2.
1."S
10
.
0 .8.6
E
2 0.4
0
.E 0.2
S00
0.8
0.8
.6
0.6
E
' 0.4
E
.4
0
S 0.2
1000
800
wavelength (nm)
00o
.S 0.2
00
.
1
8
00 .8
00 .8-
0. 6
E
2 0 .4
&0.6
E
.4
oE
0
.2
0.6
E
m .4
(U 0
C.
0 .2
800
1000
800
wavelength (nm)
00
1000
800
wavelength (nm)
-I,"S
1 .2
1.
10
1000
800
wavelength (nm)
0.2
1000
800
wavelength (nm)
R0-
1000
800
wavelength (nm)
Possible fiber coupled spectra after spectral shaping of the Ti:A120
Figure 3.3.
femtosecond laser.
76
3
After passing through the extracavity prism pair, the laser is coupled into a fiber leading to
the imaging system. Poor coupling of short wavelengths cuts some of the spectral bandwidth.
Nonetheless, quite large bandwidth can be coupled into the fiber to provide coherence gates of
less than 3 um. Possible spectra achievable using the discussed spectral shaping are displayed in
Figure 3.3. Note that the plots are individually normalized to best show spectral features. The
laser power represented by each spectrum is not the same since the spectral shaping setup causes
loss of power. For the very small bandwidth spectra shown, the power drops below 10 mW into
the fiber, which is prohibitively low for in vivo OCM imaging.
3.3.3
Fiber Broadened Femtosecond Laser Sources at 1064 nm and 1250 nrn
Methods of generating broadband light using nonlinear effects in fibers have also been
demonstrated [8, 9] and used for ultrahigh resolution optical coherence tomography [10]. These
sources take advantage of self-phase modulation resulting from confinement of an intense
electric field from a short laser pulse to generate new frequency components. In addition, fiber
broadened sources can be made compact and reliable. The femtosecond Ti:A120 3 laser discussed
above is a large tabletop laser which requires a sizeable second laser as the pump. Fiber
broadened sources require a moderate bandwidth laser to pump the fiber, which can be made
quite compact and is commercially available at some wavelengths.
Figures 3.4a and 3.4b show spectra generated from two fiber broadened femtosecond laser
sources. A spectrum generated in a germanium doped fiber pumped by a compact Cr 4*:forsterite
femtosecond laser source is shown in 3.4a. The source was developed for in vivo OCT imaging
studies by Karl Schneider, a visiting researcher at MIT. It will provide long wavelength, short
coherence length capability for OCM imaging. With over 200 nm of bandwidth and 40 mW of
fiber coupled power, the source can provide coherence lengths smaller than 5 um and may enable
cellular imaging at depths approaching 1 mm. The spectrum from this laser is used in this thesis
for broadband characterization of some components of the 1300 nm OCM imaging system but
has not yet been implemented for in vivo imaging.
Cr:Forsterite
Nd:Glass
1.2
1.2
1 -
<0.8
<0.8
0.6 -
0.6
2
E
00.4
00.4-
S
S
0.2
0.2
1100
1200
1300
1400
1500
900
wavelength (nm)
1000
1100
1200
wavelength (nm)
A
B
Figure 3.4. Spectra from fiber broadened femtosecond laser sources at 1260 nm and 1060
nm.
77
1300
Figure 3.4b shows a spectrum generated using a compact femtosecond Nd:glass laser to
pump a high numerical aperture fiber. Both the fiber and the laser source are commercially
available. The source provides over 100 nm of bandwidth centered at 1064 nm, which
corresponds to a coherence gate of around 5 um. This source will provide a future option for
high resolution OCM imaging in the 1 um wavelength range.
Interferometer
3.4
As was demonstrated in chapter 2, for fixed sample power, use of dual-balanced detection
provides enhanced sensitivity over single detector configurations. Two different dual-balanced
Michelson interferometer configurations were implemented for imaging at 1300 nm and at 800
nm. These are shown in figure 3.5a and 3.5b. At 1300 nm, an optical circulator was used to
improve the light throughput in the illumination and detection paths. Circulators are
commercially available at 1300 nm. With only 20 mW of power from the superluminescent
diode source, this component enables dual-balanced detection to be performed with the 1300
system. At 800nm, optical circulators are not available commercially. Instead two 50/50 fiber
optic couplers were used for dual-balanced detection. The Ti:A12 0 3 laser provides sufficient
power at the input that the forward loss at the first coupler can be overcome. Attenuation at the
high detector is now necessary for dual balancing as is discussed in chapter 2.
Source
Source
1
3
2
50/50
3
5051
3
Reference
2]
Reference
4
D1to
Di
Polarization
Control
50i50
2
4
Polarization
Control
50/50
2
Sample
4
Sml
D2
D2
B
A
Figure 3.5. Fiber optic Michelson Interferometer configurations used for OCM imaging.
A) 1300 nm setup using optical circulator. B) 800 nm setup using paired 50/50 couplers.
The interferometers require single-mode fiber. The 1300 nm system fibers have 9 urn core
diameter and 125 um cladding diameter with a numerical aperture of 0.11. The cutoff
wavelength is around 1260 nm. The 800 nm system fibers have 5.5 um core diameter and 125
urn cladding diameter and a numerical aperture of 0.14. The cutoff wavelength is approximately
730 nm.
3.4.1
Spectral Transmission Measurements
Flat spectral transmission curves in the circulator and the couplers are optimal for the
interferometer to avoid distortion of the spectra interfering at the detector. At 1300nm,
78
broadband fiber optics are readily available due to the demand for them from the
telecommunications market. At 800 nm, however, broadband circulators are not available and
broadband couplers are only made to custom order. Measurement of components used in the
OCM systems are provided in figures 3.6 and 3.7.
Coupler
Optical Circulator
1.2 r
-_
1:3
----- 1:4
- 4:2
3:2
1:2
2:3
I
.2
0.8
0.8-
0
0.6
I,-
0.6-
0.4
0
(0.4
0.2
0.2
1200
1300
wavelength (nm)
1100
1400
0
1000
1500
1100
1500
1400
1200
1300
wavelength (nm)
Figure 3.6. Spectral transmission measurements for optical circulator and coupler from
1300 nm system. Measurements made with Cr:Forsterite laser.
Coupler I
Coupler 2
I
-_
-1:3
- - 1:4
1:4
0.
0.8 F
4:
0.6
6 -CM
CL
0 0. 4
0
U
0.
0.4
0.2 1
0.
n
650
700
750
800
850
wavelength (nm)
900
950
650
1000
700
750
800
850
wavelength (nm)
900
950
1000
Figure 3.7. Spectral transmission measurements for couplers from 800 nm system.
Measurements made with Ti:Sapphire laser.
Transmission measurements were made using the Cr 4+:Forsterite laser at 1300 nm and the
Ti:A12 0 3 laser at 800nm. The plots display the output spectrum normalized by the input
spectrum. The large fluctuation at the edges of the spectral range is just noise in regions beyond
the source spectral bandwidth. At 1300 nm spectral transmission is relatively flat across the
source wavelength range for the coupler. The 1300 nm circulator, however, cuts off at the low
wavelengths quite sharply on the reverse path to the detector. This will limit broadband
performance with the Cr:Forsterite laser. The 800 nm couplers have overall worse transmission
79
characteristics than the couplers at 1300 nm. Nonetheless, the slope of the transmission curves is
relatively gradual and does not prohibit broadband imaging.
3.4.2
Polarization Control
Interference only results between two fields of parallel polarization. Polarization changes
that misalign the reference and sample arm electric fields can be induced by the sample or by
uncontrolled birefringence in the reference or sample arm optics or in the fibers themselves. To
optimize the interference signal against these changes, polarization paddle controllers are used in
both reference and sample paths. The paddles introduce controlled, stress-induced birefringence
to alter polarization state. Using three paddles in a A /4: 1 /2: A /4 configuration, coverage of
the entire Poincare sphere can be achieved for a monochromatic source. With broadband light,
of course, it is not possible to achieve perfect polarization control across the spectrum since a
A /4: A /2: A /4 configuration cannot be achieved at every wavelength. Nonetheless, setting the
paddles for the center wavelength provides reasonably good polarization control and
optimization of the heterodyne signal. One paddle is used in each arm to allow for calibration of
the baseline system. Looking at a sample which is not birefringent, the paddles can be set to
optimize the heterodyne signal and essentially calibrate out the reference arm and coupler
birefringence. When imaging a sample with birefringence, adjustment of the sample arm paddle
then provides compensation for induced polarization changes. If significant birefringence in the
sample arm optics exists, separation of it from sample-induced changes is difficult. Several
groups have looked with some success at polarization sensitive OCT as a mode of contrast
enhancement for standard reflectance imaging [11-13].
The diameter and number of fiber loops wound on each paddle determines its birefringence
character. The following equation determines the number of loops necessary for each paddle to
achieve A /4: /2: A/4 performance
N= 1
DL(3.1)
7rw B d2
where A0 is the center wavelength, DL is the fiber loop diameter, d, is the fiber cladding
diameter, e = 0.133, and B = 4 for A/4 operation or B =2 for A/2 operation.
3.5
Reflective Grating Phase Modulator
The heterodyne signal in OCM depends on the introduction of a phase modulation term
#(co, t). At low speed, fiber stretching phase modulators can be used to impart tiny path length
changes ( < 1 um ) to the reference arm signal, which results in a phase delay that generates the
beat signal. To provide high-speed phase modulation, electro-optic waveguide phase modulators
are available at select wavelengths common in telecommunications. These modulators, however,
do not support large bandwidth and are not available at arbitrary wavelengths. To provide for
broadband and high-speed phase modulation, a reflective grating phase delay modulator was
constructed. The grating phase modulator is similar to grating-based pulse amplitude and phase
80
shaping techniques in femtosecond optics [14-16]. The theory of operation of the grating phase
modulator is covered in section 2.6. This section emphasizes the reflective design and
characterization of its performance for broadband phase modulation.
3.5.1
Modulator Design
The reflective phase modulator design is shown schematically in top and side views in
figures 3.8. A super-achromat lens specially designed for broadband use serves as the
collimating lens. The input beam passes through dispersion compensating prisms and is
dispersed by a grating onto a curved mirror, which then focuses the diffraction orders onto the
scanning galvanometer mirror. The beam path passing through the entire system is the m = 1
order of the diffracted light. A slit filter removes higher diffraction orders which would
otherwise pass through the modulator and produce interference with the m = 1 order at the
detector. The galvanometer mirror is on a high resolution positioning stage to allow for precise
control of the offset of the beam with respect to the mirror center axis. The beam from the
galvanometer mirror reflects back to the curved mirror where the spectral components are
defocused and then recombined at the grating. A small vertical tilt of the grating introduces a
vertical displacement of the return beam from the input beam. The displaced beam is picked off
by a flat mirror and retro-reflected through the system for a second pass.
Top View
Side View
Dispersion
Prisms
compensating
Fiber
prisms
collimator
Double pass
mirror
Fiber collimator on
translation stage for
delay adjustment
.
Slit filter
Minor
Input beam
passing
mirror
over
Grating
Double pass
mirror
Curved
mirror
Scanning
Slit filter for
higher orders
Cuirdrmrro
Grating
mirror
GangCurved
Mro
Figure 3.8. Reflective grating phase delay line in top and side views.
The double-pass mirror serves two important purposes. First, it doubles the delay
characteristics of the device, allowing the modulator to deliver twice as high of a modulation
frequency for the same galvanometer scan rate. Second, the mirror helps in recoupling the beam
back to the collimator. The scanning mirror imparts an angle displacement to the return beam,
causing the beam to raster scan a line on the curved mirror and grating. The spectrally
recombined beam returning to the collimator is then displaced laterally from the input beam.
This beam walk-off effect due to scanning can create severe scan modulation on the backcoupled
reference arm power level. Modulation on the reference arm backcoupled power level shows up
directly as modulation of the heterodyne amplitude. The presence of the double-pass mirror
combats the beam walkoff effect by forcing the beam to retrace its path back to the collimator.
The input angle should be chosen to correspond to maximum grating efficiency as described
in chapter 2. To minimize astigmatism introduced by off-axis focusing at the curved mirror, the
81
grating should be as close vertically as possible to the scanning mirror. In addition, the input and
reflected beams on the curved mirror should straddle the center axis of the mirror.
To conserve reference arm power, the input lens can be set to focus at a distance much larger
than the delay line path, effectively collimating the beam through the system. Due to residual
beam walkoff, this configuration typically results in significant modulation of the backcoupled
reference arm power when the galvanometer mirror is scanning. To minimize this modulation,
the input collimating lens is typically set to focus at a distance corresponding to the path length
to the retroreflecting mirror. This position ensures that the beam is slightly defocused when
leaving and returning to the collimator, which helps with reducing scan modulation. The
reduction in modulation comes at the cost of lower backcoupled power, however.
The input collimating fiber and lens combination is mounted on a translation stage with the
axis of translation parallel to the beam. Variation of the position of the collimator changes the
reference arm path length, which adjusts the position of the coherence gate relative to the
confocal gate. Careful control of reference arm path length is required for precise alignment of
the confocal and coherence gates.
The curved mirror and galvanometer scanning mirror are mounted together on a translation
stage with the axis of translation parallel to the mirror center axis. Adjustment of the stage
position changes the distance between the grating and the curved mirror, which introduces a
second order dispersion term as described in chapter 2. Use of the phase delay line for
compensation of dispersion mismatch is common practice in OCT imaging.
To achieve pure phase modulation, the grating and mirror parameters must be chosen to
place the zero group delay offset point at a physically realizable distance on the galvanometer
mirror. Furthermore, high-speed operation requires as small of a mirror as possible to minimize
inertial load on the galvanometer. The largest mirror dimensions for galvanometers scanning at
500 Hz or faster are typically in the range of 10 - 15 mm. The grating and lens choice must
therefore be chosen to restrict the offset to below 5 - 7 mm. From expression 2.106 it is clear
that this entails relatively small focal length f and relatively low dispersion grating density
1/d. For the 1300 nm system, pure phase modulation was achieved at an offset of 2.35 mm
using a grating density of 36.152 lines per mm and focal length of 5 cm. The 800 nm modulator
achieved zero group delay at offset 3.2 mm with grating density 80 lines per mm and focal length
5 cm.
A continuous phase delay scan is generated by driving the galvanometer with a triangle wave
linear in the up and down slopes. For a given modulator configuration, the frequency and
amplitude of the drive waveform set the heterodyne frequency. Attention must be given to
choice of a heterodyne frequency that is high enough to support the expected resolution of the
system. For accurate representation of the resolution element, there must be a sufficient fringe
density under the envelope representing the structure. A demodulation frequency of 1 MHz was
achieved in the OCM delay lines using a triangle drive waveform at 500 MHz frequency with a
drive angle of about 3.6 degrees at 800 nm and 7.8 degrees at 1300 nm. Note that the heterodyne
frequency shifts with the center wavelength of the light source and must be adjusted when the
spectrum is varied. The frequency of 500 Hz is near the upper limit of the galvanometer, but the
angle drive is well below the maximum drive angle of ±20 degrees. Since imaging can be
performed on both the up and down slopes of the modulator triangle wave, a line scan rate of
1000 Hz is possible with this modulator configuration.
For imaging, the modulator drive waveform is synchronized with the fast axis of the XY
raster scan of the sample arm microscope. This provides a continuous phase modulation across
82
the image field of view. The fringe density per resolution element then depends upon the size of
the scan field. Figure 3.9 provides the fringe density for the delay line parameters used in this
thesis. For scan ranges below 250 um, the modulator still produces about 4 fringes per micron of
scan, which should be sufficient for demodulation of structures on the size scale of the focal spot.
Note, however, that for looking at submicron structure over a large field of view, the fringe count
may not be sufficient and a modulator with higher phase delay characteristics should be
designed.
Fringe Count per micron vs. scan range
inY
8
0
E
4
2
100
150
200
250
300
350
Scan Range (um)
400
450
500
Figure 3.9. Fringe count per micron for OCM phase modulators. Operation at 1 MHz
heterodyne electronic frequency is assumed.
3.5.2
Modulator Characterization
The reflective delay lines are capable of modulating the large bandwidth from the
femtosecond laser sources. Figure 3.10 shows the spectral throughput of the phase modulators.
The 800 nm and 1300 nm measurements were made with the Ti:Sapphire and Cr:Forsterite
lasers, respectively, by recording the backcoupled spectrum at the high detector of the
interferometer using an optical spectrum analyzer. Spectra were recorded with and without the
galvanometer scanning to demonstrate that the entire bandwidth is modulated.
800nm
1300nm
1.4
1.4
Input
Backcoupled
....---- Backcoupled Scanning-'
1.2
----
1
E
*-
0.8-
0.8
w-
U) 0.6
0.6 -
.
0
Backcoupled Scanning
W
E
U)
Input
-------- Backcoupled
1.2-
0.4-
o.4 0
0
Z0.2
0.2
600
700
800
900
0
1000
wavelength (nm)
1100
1300
1200
wavelength (nm)
1400
1500
Figure 3.10.
Modulator spectral throughput for 800 nm and 1300 nm systems.
Backcoupled spectra are shown with and without the modulator scanning to demonstrate
that the entire bandwidth is modulated.
83
The spectra in figure 3.10 are individually normalized to best show the spectral features. The
actual power throughput from the modulator is quite low, typically on the order of 5 %. This is
more than enough reference arm power for OCM imaging applications, though. In practice, the
reference arm return power is usually attenuated even further with neutral density filters to
prevent from saturating the photodetectors. With dual-balanced systems, the referenece arm
heterodyne gain should be set to the maximum level that does not cause voltage clipping at the
output of the transimpedance amplifier.
The order blocking slit filter is an important component of the grating phase modulator.
Because low dispersion gratings are needed to achieve zero group delay with a reasonable mirror
offset, multiple diffraction orders strike the curved mirror. If not removed from the beam path,
these orders will couple back through the system and will interfere with the m = 1 order. Figure
3.11 shows the measured reference arm backcoupled power with and without the order blocking
filter. Note that the filter together with precise alignment of the grating, curved mirror and
galvanometer mirror allow the modulation of the backcoupled power from the modulator to be
reduced to a very small level.
Without Order Block
With Order Block
33
_--
2.5
Backcoupled Power
Scan Mirror Position2P
-Backcoupled
2
2
1.5
1.5
E 0.5 -
E
0.5
-1.5
-
<
-0.5
-2
Powe
-----
-1
-0.5
0.5
0
time (ms)
1
1.5
-
-0.5 --
2
-2
-1.5
-1
-0.5
0.5
0
time (ms)
1
1.5
2
Figure 3.11. Backcoupled reference arm photodetector level with and without order
blocking filter. Sample arm is blocked during measurement.
The offset of the beam on the mirror must be precisely controlled to achieve zero group
delay. Particularly with broadband, short coherence gate systems, a group delay scan of even
several micrometers can displace the coherence gate significantly across the line scan. The
resulting heterodyne signal is then modulated with the shape of the coherence gate. A highresolution micrometer is used here to achieve zero group delay in the OCM modulators.
Figure 3.12 illustrates the heterodyne interference signals for various operating points of the
grating phase modulator. The signals were acquired using the 800 nm system after carefully
matching the reference and sample arm path lengths. These results correspond well to the
Figure 3.12a is the group delay scanning point located
simulations from chapter 2.
symmetrically to the zero group delay point. The heterodyne frequency is 1 MHz at this offset.
Figure 3.12b is the zero phase delay operating point. No heterodyne frequency results. Figure
3.12c is the zero group delay operating point. Note the continuous sinusoidal interference
fringes that result from careful adjustment of the offset position.
84
Group Delay Scanning
3.5
__3.5
A
3
Zero Group Delay
Zero Phase Delay
._._._
3
3.5
B
2.5
2.5
2-
2
2
1.5 -
1.5
1.M
1.5
0.5 -.
-0.03
0.5
5
-0.02
-0.01
0
time (ins)
0.01
0.02
C
3
2.5-
-0.02
-0.01
0
time (inS)
0.01
0.02
-0.02
-0.01
0
0.02
0.01
timf* (ins)
Figure 3.12. Operating points for grating phase modulator. Measured traces are heterodyne
interference signals recorded after the transimpedance amplifier.
3.6
Sample Arm Optics
The sample arm in an OCM system consists of a scanning confocal microscope.
Confocal
images can be obtained directly by blocking the reference arm phase modulator and recording
the DC output of the transimpedance amplifier. Several confocal microscope designs were
implemented for the work in this thesis. Benchtop microscopes were used for imaging at 800 nm
and a compact handheld probe was designed and demonstrated at 1300 nm. This section
discusses the design and characterization of these microscopes.
3.6.1
Microscope Objectives
Microscope objective lenses for investigation of in vivo OCM imaging were chosen to
correspond to the requirements of confocal microscopy. For in vivo confocal imaging, image
quality is critically dependent upon correct choice of objective lens [2]. The objective
determines the axial section thickness and lateral resolution, and is typically an important source
of loss in the optical transmission path. Table 3.1 compares the specifications of the lenses used
for work in this thesis.
Tube Length /
Focal
Working
Lens
Mag
NA
Tube Lens
Length
Distance
Immersion
Coverslip
Correction
LOMO
30x
0.9
160 mm / NA
5.3 mm
1.3 mm
Water
Yes
Zeiss
40x
0.8
Inf/ 164.5 mm
4.11 mm
3.6 mm
Water
No
Throughput @
800 nm
Price
Planachromat
-70 %
$200
Planachromat
>85%
$2,800
Table 3.1. Microscope objectives used for OCM imaging.
The lenses have medium magnification that provides a nice compromise between transverse
resolution and field of view. Both objective lenses have high numerical aperture and relatively
long working distance and require water immersion. For in vivo imaging, water immersion
microscope objectives are desirable because many tissues have an index of refraction near water.
Viable epidermis and dermis, for example, have indices of refraction of 1.34 and 1.40,
respectively. Index matching minimizes the reflection loss at the surface as well as the amount
of spherical aberration incurred upon refraction into tissue. Both lenses were initially designed
85
for use in the visible wavelength region around 500 nm. The much higher price of the Zeiss lens
is partly due to a specialized NIR anti-reflection coating which extends the 85% transmission out
past 1200 nm. The LOMO objective has no anti-reflection coating.
The LOMO objective has a finite tube length while the Zeiss lens is infinity corrected. To
account for the difference in tube length correction, the lenses should be used in slightly different
microscope configurations as discussed in chapter 2.
3.6.2
Reflective Microscope Design for Finite Tube Length Objective
A benchtop microscope based on a reflective design patented by BioRad Inc. was constructed
for use at 800 nm with the finite tube length LOMO 30X objective lens. Figure 3.13 provides an
illustration of the approximate beam path through the reflective microscope.
Curved Mirror f3
X galvo scanning
beam out of plane
of page
Raster
Lens f 4
I Plane
Objective
Mo
Collimator ff1
fiberbeam Cfibmer
Y galvo scanning
in plane of
page
Curved Mirror f2
Figure 3.13. Reflective confocal microscope design used for OCM imaging with finite tube
length objective lens.
The corresponding unfolded design and relevant notation is shown in figure 2.6a. The spot size
for the microscope is given as the fiber core size scaled by the magnifications of the two relay
pairs and the objective lens.
spot size = (fiber core)-
2
(3.2)
4
f )A,
Mo
The field of view can be determined in terms of the angles swept by the galvanometers scaled by
the magnification factors of the optics that relay the scan to the objective. The X and Y scan
durations at the sample become
= M LI
dX =
- -tan(O
4
t + 0(_
dY
dY =
0
=_M_ L,
0
L
A(3.3)
+fo ).tan (0
,+f
86
0
where L is the tube length of the objective lens and S, is the distance between the second
curved mirror
f3
and the lens
f4 .
The actual magnification of the objective is lower than the nominal magnification if the full
aperture of the lens is not filled with a uniform intensity. Achieving diffraction-limited
performance typically involves overfilling of the objective by 2-3 times, which wastes light in
the illumination path. In OCM, the axial sectioning requirement is not as stringent due to the
sectioning provided by the coherence gate. Hence, diffraction-limited performance from the
objective lens is not as critical as in confocal microscopy.
The component design parameters used for imaging with this microscope are f = 10 mm (f /
2), f 2 = f3= 50.8 mm (f / 1), and f 4 = 75 mm (f / 3). The collimator is a superachromat lens
specially designed for broadband operation at 800 nm. Spherical, gold-coated curved mirrors
provide high reflectivity at near infrared (NIR) wavelengths. The lens f 4 is a commercially
available achromat doublet designed and coated for NIR wavelengths. The illumination path
throughput was approximately 50% and the backcoupling was 45%. Throughput is measured as
the ratio of the power after the objective to the power out of the collimator.
Power After Objective
Forward Throughput = Power After Collit
Power After Collimator
(3.4)
The backcoupling is a measure of how much of the light that returns to the fiber actually couples
into it. The formula to measure backcoupling is
.
=2
x Power at Detector
Power at Sample x System Throughput
where the factor of 2 must be added to account for the 50% power loss due to the coupler. To
characterize the microscope further, the axial point spread function was measured using the
standard technique of recording the DC photodetector level while translating a mirror through
the focus [17]. Figure 3.14 demonstrates the results. The measured full width at half maximum
(FWHM) of the confocal gate is about 12 um. The transverse resolution was characterized by
imaging a United States Air Force (USAF) resolution target. The smallest pattern has a period of
4.4 um and is easily resolved with the microscope. The 10 - 90 % edge width of a scan across a
resolution element is smaller than 2 um, which corresponds to a FWHM Gaussian beam spot size
of better than 2 um as well.
The reflective nature of the design aids in eliminating some of the chromatic aberration
imparted by the lenses when used under broadband illumination. There is some astigmatism that
results from use of the mirrors off axis, however. To minimize the astigmatism, the
galvanometers should be brought as close to each other as possible, which reduces the angles at
which the beam strikes the mirror. This configuration also has the advantage of imaging both
scanners precisely to the telecentric plane in the objective. As the scanners rotate, the beam will
pivot about a single point in the telecentric plane and the focal spot will in turn raster scan a
precisely defined plane in the sample.
87
12
.
Axial Point Spread
1
>
0.80
0.64"80
2 .4
0.2-
-40
-60
-20
0
20
Axial Displacement (um)
40
60
Figure 3.14. Reflective confocal microscope characterization at 800 nm. A) OCM image
of USAF test target. The scan range in the image is 100 um x 100 um. The confocal gate is
12 um and spot size is better than 2 um. B) Measurement of confocal axial point spread
function of microscope.
3.6.3
Close-Coupled Scan Design for Infinity Corrected Objective
The infinity corrected Zeiss 40X objective lens was used at 800 nm in a benchtop microscope
design with close-coupled scanners. The optical geometry is given in figure 2.7b and redrawn
here as figure 3.15 for clarity.
Close coupled
scanners
f2
f3
Objective
Mo
Xfiber
Raster
Plane
V
Telecentric
Pupil Plane
f3
f3
f2
f,
y
f2
d
fl
Figure 3.15. Close-coupled scan design for infinity corrected objective lens
The equations for spot size and field of view with the microscope are derived similarly to
those for the reflective microsope. The expression for spot size is
spot size = (fiber core). --2.
88
(3.6)
where
f4
3
is the nominal focal length of the objective lens. Again, the actual spot size
M0
depends upon the degree to which the objective aperture is filled to achieve diffraction-limited
performance. The field of view is given by the expression
dX = dY =
IWO
(3.7)
A
where ?,, represents the scan angle swept by the X or Y scanner. The scanners are closely
situated with the center point between them located a distance f 2 away from the second lens.
The closer the scanners are to each other, the better is equation (3.7) at describing the scan field
of view. Note that the spot size and field of view depend on the same factor
which implies
f4
A3
a tradeoff between achieving small spot size and large field of view.
The implementations of the close-coupled scan design demonstrated in this thesis used a
broadband superachromat collimating lens identical to that used in the reflective microscope
design
(f;
= 10 mm, f / 2).
The relay configuration
f
2
x
f3
was varied to achieve different size
confocal gates. Figure 3.16 shows the measured variation of the confocal gate with relay lens
magnification. Measurements were made by recording the DC transimpedance voltage output as
a function of position when translating a mirror through the focus. A 1 um resolution
piezoelectric stage was used.
All lenses used in the relay configurations were NIR achromat doublets of 1 inch
diameter except the 164.5 mm focal length tube lens, which had a diameter of 35 mm. For all
relay configurations, the illumination path throughput and detection path backcoupling was high
for a typical confocal microscope operating at near infrared wavelengths. The measured values
for all relay configurations are listed in table 3.2. All measurements were made at 800 nm using
the modelocked Ti:Sapphire laser with greater than 100 nm bandwidth.
The throughput values decrease with increasing magnification of the relay lens pair because
the overfilling of the objective increases with larger beam diameter. The loss at the objective can
be compensated by increasing source power to provide fixed sample power. The backcoupling
values are normalized to the throughput and therefore essentially provide a measure of the
aberration in the system. Backcoupling is extremely sensitive to alignment of the mirror in the
sample, however. Care was taken to optimize alignment before recording each measurement.
89
Confo cal Gate
1.2
1
0)
-J
0.8
A
0
U
0)
a, 0.6
0
U)
a)
B
0.4
C
0.2
-
D
E
n
-40
-30
-20
20
10
-10
0
Axial Displacement (um)
30
40
Figure 3.16. Confocal axial point spread function for various relay configurations f2
x
f3 .
A) 100 x 75 mm, FWHM ~ 66 urn B) 100 x 100 mm, FWHM ~ 30 um C) 100 x 164.5
mm,FWHM - 12 um D) 75 x 164.5 mm, FWHM - 7 um E) 50 x 164.5 mm, FWHM ~
3um
Relay
50 x 164 mm
75 x 164 mm
100 x 164 mm
100 x 100 mm
100 x 75 mm
Throughput
56.50%
66.90%
67.20%
72.20%
74.50%
Backcoupling
72.60%
69.80%
74.90%
68.80%
72.70%
Configuration
Forward
Table 3.2. Illumination path throughput and detection path backcoupling for various relay
configurations.
The various microscope configurations allow high lateral resolution with transverse focal
spot sizes measuring below 2 um. The 50 x 164 configuration provided a Gaussian spot size of
better than 1 um, as measured with the 10 - 90 % edge width technique. Even the 100 x 100
relay setup still provided spot size below 2 um with 30 um confocal gate. These results
empirically confirm that the transverse resolution can be maintained without extremely high NA,
potentially allowing for OCM imaging at the cellular level with relaxed confocal gate. Figure
3.17 illustrates the resolving power of the microscope for both very high NA and relaxed NA.
The smallest element of the USAF test target is clearly visible in both confocal images.
90
F7
-w
B
A
ga
igauees3.17.
n
nuerial apere.A
ui
manflimes of USAF teestcatwihdfeentri
aTsnge physticalhspangenteae, the bgamvanoete
s ab0u
irroscin btoubloe priularlynth
Y direction about another point. Using the lens equation, an expression can be derived for the
spacing between the X and Y pivot points.
Pivot Error
=
-
r[f+
)f(f±
M
)/M
(3.8)
where dM is the actual mirror spacing. The physical spacing between the scanners was slightly
less than 6 mm. When the relay pair is set with a short lens f2 = 50.8 mm followed by the tube
lens fL = 164.5 mm, this spacing dM is magnified to over 30 mm spacing around the telecentric
plane. With such significant pivot error, the beam clips on the limiting aperture of the objective
at the extreme angles of the X and Y scans. Loss of signal at the edges of the image results.
Similarly, misplacement of the objective with respect to the tube lens moves the telecentric pupil
plane with respect to the images of the scanners and can result in beam clipping as well. Figure
3.18 demonstrates the effect of imperfect imaging of the scanners to the pupil plane. The data
was recorded in confocal mode at the output of the transimpedance amplifier with the reference
arm blocked. Figure A shows the measured backcoupled detector signal for a single line scan of
the image and figure B shows the resulting loss of signal at the image fringes. Despite the beam
91
clipping, note the clarity of tiny details on the mirror surface provided by the high-resolution
configuration.
To minimize vignetting of the beam at the exit pupil of the objective lens, the scanners
should be set as close to each other as possible. Correct choice of focal lengths f2 and f3 can
minimize the pivot error as well, but this generally comes with a loss of relay magnification. In
confocal microscopy applications where small axial section thickness is critical, this tradeoff is
significant. In order to preserve confocal axial section thickness, the relay magnification must be
high enough to overfill the objective. The degree of clipping of the beam must also be
minimized to provide crisp images over a large scan field. Careful design of the objective lens
and choice of relay lens parameters can be used to combat the problem, but often the less
complex solution for confocal microscopists involves using a design with the scanners separated.
X Scan Modulation
r
- Bakcoupled Power
B can Mirror Position
0.5-50
-4
.
.0
-1
0
10
20
30
field of view in the image is greater than 100 ur x 100 um.
3.6.4
Performance Under Broadband Illumination
To evaluate the effect of broadband illumination, the confocal point spread function was
measured and compared with the Ti:A12 0 3 operating in continuous wave (CW) and in
modelocked (ML) state. Measurements were made with the reflective microscope with over 140
nm of bandwidth in the modelocked case. As shown in figure 3.19, the point spread function is
only slightly broader under broadband illumination compared to monochromatic illumination.
The CW trace shows oscillations characteristic of uncompensated spherical aberration in the
objective lens. The absence of oscillations in the point spread taken under broadband
illumination is likely due to the presence of chromatic aberration. The chromatic aberration
results in broadening of the point spread as well as averaging out of sharp features that would
appear in any single frequency point spread.
92
Confocal Gate
1. 9
.------- ML
-- CW
1-
-
Z
0. 8
0
0. 6
a,
Z
1%
0. 4
0. 2
0i
-40
.
-20
0
20
40
Axial Displacement (um)
Figure 3.19. Confocal axial point spread function for monochromatic (CW) illumination
and broadband (ML) illumination with 140 nm bandwidth.
3.6.5
Combined Gating Effects
Various operating regimes for short coherence gate imaging were qualitatively investigated
by comparing OCM and confocal images. Figure 3.20 provides OCM images for comparison
with confocal images provided in figure 3.17.
Figure 3.20. OCM images demonstrating effect of combined confocal and coherence
gating. A) Coherence gate - 3 um / Confocal gate - 3 um , FOV = 130 x 130 um B)
Coherence gate ~ 3 um / Confocal gate ~ 30 um, FOV = 190 x 190 um.
93
Figure 3.20a demonstrates the short coherence gate, short confocal gate operating regime. Due
to small tilt in the mirror, zero path difference between reference and sample arms can only be
matched over a small portion of the image. When using short coherence and short confocal gate,
the heterodyne signal is extremely sensitive to precise overlap of the two gates. The signal is
completely lost for small delay mismatches. By contrast, figure 3.20b demonstrates the
interaction of a short coherence gate with a relatively long confocal gate. In this case, the gates
are less sensitive to small path mismatch and the heterodyne signal can be maintained across the
entire field of view. These results highlight the fact that the extreme limit of short coherence and
short confocal gate is impractical for in vivo imaging unless a mechanism for accurate, real time
control of path mismatch is devised. The long confocal gate combined with a short coherence
gate, however, provides a suitable working regime for investigating OCM for in vivo imaging.
3.6.6
Compact Handheld Imaging Probe
A compact, handheld microscope was developed to facilitate the transition of OCM
technology to clinical applications. The probe is shown schematically in figure 3.21. MIT
graduate student Pei-Lin Hsiung originally designed the probe for low numerical aperture
imaging and it was subsequently modified here for microscopy. The probe uses the closecoupled scan design discussed in section 3.6.3. A fiber input is collimated by a lens of focal
length 11 mm. The beam strikes a pair of closely spaced, compact galvanometers that impart
angles in orthogonal planes. A pair of lenses with f 2 = 25 mm and f, = 40 mm relays the angle
scan to a 30 X, 0.9 NA objective lens. The entire probe measures about 15 cm in length.
fiber
f2f
f2 + f3
f3
Objective
y-galvo
2
3
30x 0.9 NA
x-galvo
Figure 3.21. Handheld microscope for in vivo OCM applications.
The handheld probe was characterized using the superluminescent diode laser source at 1300
nm with 65 nm bandwidth. A confocal gate of 15.5 um and transverse resolution of better than 3
94
urn were achieved with a field of view of better than 100 x 100 um. Figure 3.22 presents the
results of characterization.
Confocal Gate
1.2r,
I
*1;
e 0.8
~1
0
(J
e 0.6
0
0
0
.2 0.4
e
0.2}
-80
-60
-40
-20
0
20
40
Axial Dis placement (urm)
60
80
Figure 3.22. Characterization of handheld microscope. A) OCM image of USAF
resolution target. Field of view = 145 um x 175 um. B) Measurement of confocal axial
point spread function.
3.7
Receiver Specifications
Detection of the optical heterodyne interference signal was performed with high-speed, lownoise receivers initially designed for use in OCT systems. The basic block diagram and
description of a receiver for low coherence interferometry is provided in section 2.4.4. A pair of
dual balanced photodiodes converts the incident optical interference signals into electrical
currents. The currents are added at the input node to a transimpedance amplifier and the sum
current is converted to a voltage via the feedback transimpedance resistance. Filtering of the
electrical signal removes wideband noise outside of the frequency spectrum of the image signal.
The filter center frequency and bandwidth are set to match the center frequency and bandwidth
of the heterodyne signal. Log demodulation is then performed to detect the envelope of the
signal and to compress the large dynamic range OCT signal into a dynamic range that can be
sampled accurately by the data acquisition card and subsequently displayed in an image. The
specifications for the receivers used for this thesis are supplied in table 3.3. The 1300 nm
receiver was designed by Eric Swanson and the 800 nm system was designed by Ingmar Hartl.
Transimpedance
Filter Center
Filter
Filter
Wavelength
Type
Responsivity
Resistance
Frequency
Bandwidth
Order
Demodulation
1300 nm
800 nm
InGaAs
0.75 A/W
0.51 AW
100 kOhm
744 kOhm
900 kHz
1.0 MHz
550 kHz
170 kHz
2nd
3rd
Log
Log
Diode
Si
Table 3.3. Specifications for OCM receivers used at 800 nm and 1300 nm.
95
Choice of diode type is dictated by the quantum efficiencies of the material at the wavelength
of interest. Silicon works well in the wavelength range around 800 nm but drops off
dramatically above 1000 nm and cannot therefore be used for applications at 1300 nm. At 1300
nm, however, InGaAs diodes provide high quantum efficiency. Note that the higher responsivity
of InGaAs at 1300 nm provides a sensitivity advantage for the 1300 nm receiver compared to the
lower responsivity of Si at 800 nm.
The filters are low order passive Butterworth filters. The required filter bandwidth is
determined by the size of the resolution element and the velocity of the image line scan. An
approximate expression for the bandwidth can be written as
1
Ax
1 FOV
*AX
(3.9)
TSCAN
where Ax is the size of the resolution element, v, is the scan velocity, FOV is the length of the
field of view, and TSCAN is the acquisition time for a line of the image. For a triangle wave with
acquisition on both slopes, the acquisition time is the inverse of twice the waveform frequency.
Note that the bandwidths of the filters used in the specified receivers are quite large to
correspond to the relatively high group delay scan velocities used in OCT systems. High
resolution, high speed OCT typically seeks to maintain an axial resolution of several
micrometers over a group delay scan of 1-3 mm. In OCM, however, the field of view is
significantly smaller. Hence for similar resolution element size, the required filter bandwidth is
lower.
Use of a log demodulator may not be optimal for OCM imaging. OCT systems use a log
demodulator in order to compress the large dynamic range of the heterodyne signal into a
representable image. Large reflections from the surface must be shown together with very small
reflections from deeper layers. In OCM, however, the en face image should have significantly
less dynamic range than cross-sectional images. Use of a log demodulator for a small dynamic
range image likely reduces contrast. The optimal processing scheme for OCM may be a linear
demodulation technique with variable gain adjustment. Future work will investigate this option.
3.8
Image Acquisition and Processing
The image acquisition scheme is a modification of a flexible high-speed OCT acquisition
system designed by Tony Ko [18]. The hardware setup is shown schematically in figure 3.23. A
personal computer (PC) is configured with two 12-bit data acquisition cards. The first card is a
multi-function device with both analog to digital (A/D) and digital to analog (D/A) capability. A
second D/A card is added to provide additional output capacity. The analog to digital (A/D)
converter operates with 5 MHz maximum sampling rate while the digital to analog (D/A)
converter can use up to 2 MHz update rate. Analog outputs provide drive signals to the
galvanometer controller cards, which then supply feedback position control to the scanners. A
function generator is used to generate the fast axis drive waveform. The function generator is
controlled by the software interface via GPIB and triggered with a timing signal from the multifunction A/D card. Fiber optic inputs from the interferometer feed photodiodes in the electronic
96
receiver. The demodulated output from the receiver is then sampled by the A/D card. An image
is created in real time using a windows based software interface.
A/D
Receiver
SgaInElectronic
Demod
Out
D1
D/A
Modulator
Drive
Fn Gen
PCX
Galvo
D2
CFiber
Inputs from
Photodetectors
Modulator
Controller
Slow X Scan
To Modulator
Galvo
To Microscope
Controller
GPIB
Function
Generator
Y Galvo
Drive
Fast Y Scan
Controller
X Galvo
_
_
To Microscope
Y Galvo
Figure 3.23. Schematic of acquisition and control hardware.
3.8.1
Timing and Synchronization
The recorded timing signals for the OCM system are provided in figure 3.24. The triangle
modulator drive results in a linear phase ramp and continuous interference fringes over the
duration of the up and down ramps. The position of the galvanometer mirror generating the fast
axis (Y) of the microscope raster scan is synchronized to the modulator scanner position output.
Two timing pulses are generated on an analog output channel for each period of the triangle
drive waveform, one each for the up and down slopes of the triangle. Each timing pulse triggers
the generation of a Gate signal which allows digitization of the input heterodyne signal to
commence. The duration of the Gate signal is chosen to be an integer number of cycles of the
underlying pixel clock to ensure that the same number of points are acquired for each line of the
image. The Gate signals are positioned symmetrically on the most linear portion of the
modulator scanner position waveform to provide the most constant heterodyne signal. The Gate
signal also serves as the trigger signal for the function generator to output another period of the
fast Y axis drive waveform. At high speed, images are acquired on both the up and down slopes
of the triangle. To allow the software to distinguish between up and down slopes so that an
upright image is always displayed, a 'Start' signal is generated and used to trigger the
acquisition. The Start signal is triggered by the first of the two timing pulses. When high, the
Start signal allows acquisition to occur, thereby eliminating the 50% chance that acquisition
starts on the wrong slope of the triangle and produces an inverted image. The slow axis drive
waveform (not shown) is updated on the falling edge of each Gate signal. The number of update
points chosen for the slow axis waveform thereby sets the imaging frame rate. For example, a
500 Hz triangle drive for the modulator results in 1000 line scans per second, or 1 ms per line.
97
Choosing 500 pixels across the image in the slow axis then sets the frame rate at 2 images per
second. Choosing 250 pixels provides 4 frames per second.
Acquisition Timing Signals
41--
Heterodyne Signal
0
-2
-1.5
-1
.2
-1.5
-1-0.5
-0.5
0
0.5
1
0
0.5
1
10
2
-101
-2
1011
-1-
1
-1.5
-1
1
1
-0.5
0
Gate
0.5
I
1
--- Start
0-10
-2
-1.5
-1
-0.5
0
10
10
-2
--
-1.5
-1
-0.5
time (ms)
0
1
0.5
0.5
T1min
Pulse
1
Figure 3.24. Acquisition timing signals for high-speed OCM system.
3.8.2
Software Interface
The Windows-based software interface used for the OCM system was created by Tony Ko
for high speed OCT [18]. Modifications were made in C++ to incorporate the control signals
and timing synchronization for the microscope scanning mirrors and the modulator scanning
mirror. The software used a double-buffered acquisition scheme to allow for real time
acquisition. With double buffering, the system memory holding incoming digitized data is
configured as a circular buffer consisting of two half buffers. When the first buffer is filled with
data, it can be processed for display while the A/D card continues to fill up the second buffer.
The effectiveness of the double-buffered scheme relies on the ability of the computer to process
the data in the first buffer before the second buffer is full and writing begins again in the first
buffer. To avoid from having to turn off and restart the digitization repeatedly for each line scan,
a method called scan clock gating is used. With scan clock gating, the internal digitization clock
of the acquisition card is gated by a TTL signal. This Gate signal is shown in figure 3.24. In this
way, the acquisition is configured to pause when the Gate is low and the need for continually
restarting acquisition is eliminated. The real-time display of images to the screen is performed
using DirectX technology.
98
3.8.3
Zipper effects
Acquisition of images on both the up and down slopes of the triangle wave often leads to
image distortions typically known as 'zipper'. Zipper results from inexact pixel correspondence
between line scans on the up and down slopes and can result from a number of sources. First,
zipper can result from improper placement of the Gate signals with respect to the modulator
position waveform. The gates should be symmetrically located on the up and down slopes so
that the scanner positions represented by pixels in the up and down Gates correspond.
Nonlinearity in the velocity on the up or down slope can also cause a lack of correspondence
between pixels. This is particularly noticeable around the turnaround points of the galvanometer
triangle drive waveform. This zipper effect can be minimized by reducing the overall size of the
acquisition gate and driving the galvanometer harder to achieve the same effective scan field. A
second form of zipper results in the OCM images due to inexact positioning of the beam offset at
the zero group delay position. If the beam is slightly displaced from the zero group delay
position, then the envelope of the coherence gate can show up as a modulation of the image
intensity across the line scan. This is particularly relevant when the coherence gate is very small.
Differences between the group delay scans of the up and down slopes can result in slight
displacement of the coherence envelope modulation, resulting in what appears to be line to line
image intensity modulation. High precision, linear galvanometers and careful positioning of the
beam to zero group delay position can help eliminate this problem. If necessary, the effect can
be eliminated completely by using only one side of the triangle wave for acquisition. This,
however, sacrifices speed. Use of faster, resonant galvanometers with acquisition on only one
side may provide a solution to the problem.
Another form of zipper results from inexact correspondance between the fast Y scanner
position and the modulator position. If the two position waveforms are displaced or if the
linearity of the two waveforms differs, then the Gate signals may not be symmetrically located
on both the modulator and the Y-axis waveforms. Care must be taken to synchronize the
modulator and Y-axis waveforms as closely as possible to minimize this effect.
3.8.4
Sampling Criterion
The pixel density in the images must be chosen to obey the Nyquist criterion. The Nyquist
criterion states that no loss of information occurs if sampling is greater than two times the
frequency bandwidth of the signal. In reality, perfect reconstruction of the original signal from
the sampled data requires an ideal low-pass filter. As these are unavailable, Webb suggests that
2.3 times the signal bandwidth should be taken as a minimum practical sampling frequency [19].
For the OCM images presented in this thesis, pixel density at 4 frames per second is 1375 x 250.
The images are significantly oversampled in the fast raster scan direction. Sampling is limited,
however, in the slow scan direction because lower pixel density is needed to achieve higher
image frame rates. Typical field of view in the slow axis dimension is 100 - 150 um, which
results in better than 1 pixel per micrometer. Since cells are typically larger than 5 - 8 um, this
sampling rate is sufficient for resolving cellular features. However, for imaging micron and
99
submicron sized structures, including subcellular structures, this is likely near the Nyquist limit
and should be improved in future system upgrades.
3.9
System Sensitivity Measurement
System sensitivity is measured looking at a mirror in the focus of the microscope. The
modulator is set to operate in OCT mode with the beam offset on the galvanometer mirror at the
1 MHz group delay scanning point located symmetric to the zero group delay point with respect
to the mirror axis. The microscope galvanometers are not scanned. The resulting line scan
traces the axial point spread function corresponding to the coherence gate. Sensitivity is
determined by progressively attenuating the sample arm light using calibrated neutral density
filters until the signal is no longer detectable. Insertion of a 3.0 OD (optical density) filter in the
sample beam transmission path, for example, results in an attenuation of 60 dB in optical power
since the attenuation occurs in both forward and reflected light paths.
In order to obtain maximum sensitivity and axial resolution, dispersion in the reference and
sample paths must be matched accurately. An analysis of the effects of dispersion on the OCT
point spread function is provided in section 2.4.3. Recall that unbalanced dispersion results in a
broadening of the point spread, chirping of the heterodyne frequency, and reduction of the
heterodyne amplitude. To achieve dispersion matching in the OCM system, the presence of
large amounts of glass in the microscope optics must be balanced by adding glass to reference
arm path. The glass in microscope objective in particular presents a problem because the types
of glass and thicknesses of the lenses are typically proprietary information. Hence, one must use
an approximate matching technique. In the work for this thesis, two adjustable prism pairs of
glass types SFL6 and LakN22 were positioned after the collimator in the modulator as shown in
figure 3.8. These glass types were chosen to correspond exactly to the glass in the achromat
lenses used for the intermediate relay optics. Because the collimating lenses in the reference and
sample arms are the same, only the dispersion produced by the microscope objective remained.
Despite lack of exact knowledge of the objective glass types, the dispersion could be reasonably
well compensated by adding further SFL6 and LakN22 to the reference path. SFL6 is much
more dispersive than LakN22, and together the two glass types provided a reasonable coarse and
fine adjustment. The degree of dispersion matching was quantified by measuring the heterodyne
amplitude and the width of the point spread function. A relative measure of the heterodyne
amplitude is given by the fringe contrast of the detector signal, defined as the ratio of the
measured amplitude of the oscillating heterodyne signal to the predicted amplitude.
Measured Amplitude _VPk /2
Fringe Contrast = Prediced Amplitude - VRV/
Predicted Amplitude
(3.10)
2 -V ,Vs
represents the measured peak-to-peak amplitude of the oscillating signal and VR and
Vs are the DC levels of the reference and sample arm detector powers, all measured after the
transimpedance amplifier. The width of the point-spread function can be compared roughly to
the predicted width from equation (2.67).
The effects of dispersion are shown qualitatively in the measured axial scans displayed in
figure 3.25. Figure 3.25a results from excess dispersion in the sample while figure 3.25c is
Here,
VPPk
100
generated by excess dispersion in the reference arm. In both cases, the point spread is broadened
and the heterodyne amplitude reduced compared to the matched dispersion trace in figure 3.25b.
Note also that chirp is present in both 3.25a and 3.25c, although the frequency chirp is in the
opposite direction for the two traces.
- 0.78 mm SFL6
Dispersion Matched
1.4
1.4
1.2 --
1.2
,,
+ 0.78 mm SFL6
,1.4
,
1.2-
--
0.4 -<
0.2 --
<0.4 -<
0.2
,
--
-3
-2
A
.1
0
1
time (ms)
2
3
-3
0.4-0.2 --2
-1
0
time (ms)
B
1
2
3
-3
-2
-1
0
time (ms)
1
2
3
C
Figure 3.25. Effects of dispersion on axial point spread function. A) Excess dispersion in
the sample arm. B) Matched dispersion. C) Excess dispersion in the reference arm.
With careful dispersion balancing, fringe contrast between 85% and 100% was achievable,
depending on the bandwidth and the center wavelength of the source. Dispersion mismatch is
most deleterious for lower NIR wavelengths and for large bandwidths. At 800 nm with over 100
urn bandwidth, the system sensitivity will suffer dramatically if dispersion is not matched
accurately. At 1300 nm with 65 nm bandwidth, however, dispersion mismatch is less critical.
The origin of the wavelength dependence in dispersion characteristics is the wavelength
dependence of the index of refraction of the glasses used. The index variation is steeper around
the lower NIR wavelengths than the higher wavelengths.
In the shot noise limit, it was demonstrated in section 2.4.6 that the system sensitivity does
not depend on the reference arm power. In the presence of excess noise sources, however, the
reference arm power level does have an effect on the signal to noise ratio [20]. From section
2.4.5 it is seen that the electronic receiver noise does not depend on the optical power incident on
the photodiodes. The signal to receiver noise ratio therefore increases with increasing reference
power. For a single detector, however, the excess intensity noise from the laser source also
increases with increased reference power, and the resulting signal to excess noise ratio decreases.
The crossing point between the rise in signal to receiver noise ratio and the fall in signal to
excess noise ratio defines an optimal reference arm power for the single detector configuration.
Using dual balanced detection, the excess noise can be largely cancelled, leaving receiver noise
and shot noise. Hence, in the dual balanced configuration, the signal to noise ratio should
increase with increasing reference arm power. In practice, the reference arm power for a dual
balanced receiver should be as high as possible without saturating the heterodyne signal at one of
the receiver stages.
Sensitivity measurements were made with the reference and sample power levels used for
imaging. Typical incident power on the sample is 4 - 6 mW, and power returning from the
reference arm was generally held at between 10 and 100 uW. The signal to noise ratio was
determined in two ways. First, the sample arm attenuation was increased until the log
demodulated signal was no longer visible on the image. This method has the advantage of
101
providing the truest measure of sensitivity as is seen in the OCM images. Using this method, the
SNR for the reflective microscope design was measured to be over 85 dB at 800 nm. The higher
throughput of the Zeiss objective along with careful alignment led to a sensitivity of better than
90 dB for the close-coupled scan design microscope at 800 nm. For the handheld probe at 1300
nm, the SNR was recorded at around 70 dB. The lower number is likely due to lower throughput
in the probe at 1300 nm and also to difficulty in the measurement. The encased probe does not
easily allow for insertion of attenuating ND filters and some misalignment likely results in the
measurement process.
The technique of measuring sensitivity from the image on the screen has the disadvantage of
being difficult to quantify a precise number for the SNR. The filters come in discrete values and
fine tuning of the attenuation is difficult. For systematically determining a precise value,
calculation of the SNR is made from the interferometric fringe signal taken after the filter and
before the log demodulator. The fringe signal is acquired for a set attenuation value A. The
power in the heterodyne signal is taken as the square of the voltage amplitude of the oscillating
heterodyne signal and the power in the noise is determined by numerically computing the
variance from the recorded trace. The formula for signal-to-noise ratio is given as
K
SNR = A+10log
V2
(3.11)
pk
var[n(t)])
Typically, three or more distinct measurements are averaged together to give a more accurate
value for the sensitivity. SNR values for the close-coupled microscope design at 800 nm using
this calculation method are reported in table 3.4. Attenuation in the sample arm was 60 dB. For
comparison, sensitivity measurements for several different reference arm power levels are
included as well as for a single detector configuration. Note that the SNR improves with
increased reference arm power and with dual balanced versus single detector configuration. The
improvement in performance in the presence of noise sources other than shot noise is responsible
for the larger sensitivity values.
Configuration
Power Incident on
Sample
Backcoupled
Reference Power
Sensitivity
Dual Balanced
Dual Balanced
Dual Balanced
6.0 mW
6.0 mW
6.0 mW
16 uW
30 uW
106 uW
93.12 dB
94.87 dB
95.68 dB
Single Detector
6.0 mW
30 uW
91.30 dB
Table 3.4. Calculated sensitivity values for various system parameters. The close coupled
microscope was used at 800 nm with dispersion carefully matched. The relay lens
combination was 100 x 100 mm. Attenuation in the sample arm was 60 dB.
3.10
Axial Resolution Measurements
The axial sectioning capability or axial resolution of the OCM system is determined by the
combination of the coherence and confocal gates. The confocal gate measurements were
102
presented in section 3.6. The coherence gate was measured by acquiring an OCT axial scan of a
mirror in the sample. To most accurately measure the coherence gate, the numerical aperture of
the sample arm optics should be dramatically reduced so that the confocal gate has very little
effect. Since this involves aperturing the beam or changing the final lens, a simpler more, more
approximate method was used here. The mirror was moved out of the focus of the microscope to
a region where the light level from the sample remained approximately constant over the range
of the axial scan. The modulator was then set to the 1 MHz group delay scanning point and the
OCT trace recorded. The time axis of the axial scan was converted to position by measuring the
length of the group delay scan using the reference arm path adjustment micrometer. Figure 3.26
presents the typical measured coherence gate used for in vivo imaging. The light source
bandwidth was better than 120 nm and the FWHM of the coherence gate measures less than
three micrometers, which should be sufficient for cellular imaging in scattering tissue.
Coherence Gate
3
2.5-
FWHM<3um.
2-
.5
E
1 0.5-
0
-15
-10
-5
0
delay (um)
5
10
15
Figure 3.26. Typical measured coherence gate for 800 nm system. The source bandwidth
was better than 120 nm to provide a coherence gate of less than 3 micrometers.
The combined coherence and confocal gate was also measured for the 100 x 100 mm relay
configuration. The confocal gate in this case was ~ 30 um and the coherence gate was better
than 3 um. The measurement was taken by recording the heterodyne fringe signal as a mirror
was translated through the focus of the microscope. The recorded fringes were demodulated and
the amplitude was then plotted as a function of mirror translation. The trace was measured with
30 dB attenuation in the sample arm and scaled according to the measured sensitivity value.
Figure 3.27 displays the results. The combined gate has FWHM width less than 3 um. The
smaller peaks located outside of the main peak are likely spurious reflections somewhere in the
sample or reference arm optics. They could result from elements in the microscope objective or
from multiple reflections at the fiber connections. They should be carefully tracked down and
eliminated for optimal system performance.
103
Confocal + Coherence Gating
0
-20-
-40-
-60-
-80-
-100
-200
100
0
-100
axial displacement (um)
200
Measurement of combined confocal and coherence gate.
Figure 3.27.
separately, the confocal gate was 30 um and the coherence gate was 3 um.
104
Measured
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105
106
Chapter 4
In Vivo Imaging Results
4.1
Overview
The high-speed, broadband OCM system discussed in chapter 3 was applied for in vivo
imaging of both animal and human tissues. This chapter presents and discusses the results of in
vivo imaging at 800 nm with short coherence gate. Section 4.2 discusses imaging of the Xenopus
laevis tadpole, an important animal model for studies in developmental biology. Section 4.3
describes the potential for imaging cellular structure in human skin. In particular, the potential
of using enhanced axial sectioning provided by a short coherence gate to relax the requirement
for high numerical aperture optics is demonstrated. Finally, section 4.4 discusses initial results
using a handheld imaging probe at 1300 nm.
4.2
Imaging of an Animal Model: Xenopus laevis Tadpole
Studies in developmental biology seek to understand the genetic and molecular processes that
create an entire organism from a single cell. Small, easily handled model organisms with rapid
reproductive cycles are generally chosen for study with the hope that fundamental mechanisms
of development are conserved across species. These models include amphibians such as
Xenopus laevis and Rana pipiens, fish such as Brachydanio rerio, and insects such as Drosophila
melanogaster. To monitor development of such organisms at the cellular level, a particular need
exists for in vivo imaging techniques capable of visualizing tissue microstructure at consecutive
time points.
Optical coherence tomography (OCT) has shown exciting promise for this
application [1-4]. OCM, as a high-resolution extension of OCT, offers further capability of
investigating embryonic morphology.
As a demonstration of the potential for OCM in developmental biology, Xenopus laevis
tadpole was imaged. Figure 4.1 provides in vivo images of Xenopus cellular and tissue
morphology. The images were acquired at 800 nm with the close-coupled scanner microscope
design. The confocal gate was 30 um and the coherence gate measured approximately 3 um.
Incident sample power was 5.5 mW resulting in a sensitivity of better than 94 dB. Images were
acquired at 4 frames per second with a field of view of 210 um x 144 um. Image processing
consisted of adjustment of the aspect ratio and the gray scale. The tadpoles were imaged under
anesthesia and subsequently sacrificed in accordance with approved protocol on file with the
MIT Committee on Animal Care (CAC).
Large mesenchymal cells in the dorsal spinal cord region of the tadpole are clearly visible.
The cells measure 20 - 30 urn in diameter. Cell nuclei and cell membranes are also
distinguishable. The nuclei measure less than 10 urn in diameter. In addition, cells bordering the
outside world are visible in figures 4.1g and 4.1h. These are likely epithelial cells.
107
Figure 4.1. In vivo cellular images of Xenopus laevis tadpole. Large mesenchymal cells
are visible in the dorsal spinal cord region of the tadpole. Nuclei (N), cell membranes (CM),
and epithelial cells (EC) are clearly distinguishable. Image field of view is 210 um x 144
um.
108
Figure 4.2. In vivo images of blood flow in Xenopus laevis tadpole.
Vessels of varying
sizes are distinguishable along with individual blood cells (BC) suspended in the vessel
lumens. Image field of view is 210 um x 144 um.
109
Real time imaging capability was demonstrated in the Xenopus tadpole by imaging blood
flow in vessels of varying sizes. Figure 4.2 provides still frame images of blood cells in such
vessels. Video sequences of flow can easily be reconstructed from saved frames or can be
recorded directly with an S-VHS recorder. Note in some of the blood flow images that the blood
cells are not clearly resolved. These cells may be passing through the plane of image at high
flow velocity. Imaging at greater than 4 frames per second is desirable to clearly resolve the
faster moving cells. In figure 4.2a and 4.2b, for instance, the slower moving cells along the
vessel wall are resolved while the faster moving cells in the center are not clearly resolved.
Future improvements will include increases in imaging speed.
4.3
In Vivo Imaging of Human Skin
Imaging of Xenopus tissue does not accurately assess the potential of OCM for imaging of
many human tissues of interest. Human tissue is generally much more highly scattering than that
of Xenopus and the cells are typically not as large. To assess capability of OCM for applications
in clinical medicine, the technique must be demonstrated in a representative human tissue. Skin
is chosen here for a number of reasons. First, skin is the most easily accessible tissue, allowing
imaging with relatively large prototype benchtop microscopes. Second, skin has a distinct
layered structure that can be recognized in images. Third, skin is among the most difficult of
tissues in which to image due to its heterogeneity and therefore provides a rigorous test of the
imaging modality. Finally, skin imaging has many relevant applications in clinical medicine and
has been approached with other modalities, including OCT and confocal microscopy [5-10]. The
growing body of work looking at both normal and pathologic skin with laser scanning confocal
microscopy in particular provides a suitable benchmark with which to judge the performance of
the OCM system [11-20].
Figure 4.3 demonstrates the layered structure of skin in a histologic section stained with
hematoxylin and eosin (H & E). The epidermis layer of skin is organized as a keratinizing,
stratified squamous epithelium. The epithelial cells, known as keratinocytes, are organized into
basal, spinous, granular, and cornified layers that correspond to progressive stages of
differentiation. The basal cell layer consists largely of mitotically active keratinocytes with
pigmented melanocytes intermixed. Newly generated keratinocytes from the basal layer will
follow a life cycle in which they progressive from immature basal cells to non-viable, terminally
differentiated corneocytes in the stratum comeum. The spinous and granular cell layers then
represent intermediate stages of differentiation. The entire life cycle from basal cell to
comeocyte typically occurs in around 14 days [21]. The underlying support layer to the
epidermis is called the dermis. It consists of a mixture of connective tissue, blood vessels, and
duct structures important for the physiologic function of the organ.
From an optical imaging perspective, the skin presents a number of index discontinuities and
irregular surfaces that induce aberration and can dramatically degrade the confocal axial point
spread function. In vivo confocal microscopy in human skin requires careful attention to choice
of objective lens and immersion fluid to obtain high contrast images [10]. The enhanced axial
sectioning provided by OCM may help to relax the need for such high-quality optics in the
probe, thereby enabling cellular imaging with probe designs not possible for use with confocal
microscopy.
110
Figure 4.3. Stained histologic section of human epidermis. The stratified epidermis
consists of basal (B), spinous (S), granular (G), and stratum comeum (SC) cell layers. The
underlying papillary dermis contains blood vessels and duct structures important for
physiologic function of the skin. Image excerpted from Freinkel & Woodley [21]. Scale
bar is approximate.
4.3.1
Exposure Limits for Microscopy
To image with maximum sensitivity, it is desirable to operate with incident sample power
near the maximum safe exposure level. The American National Standards Institute (ANSI)
produces guidelines for determining the maximum exposure assuming illumination over a large
effective aperture. In scanning optical microscopy, however, a tiny focal spot is scanned rapidly
over the tissue, creating significantly different exposure conditions. Furthermore, use of
modelocked laser sources with short pulse illumination is not properly accounted for in the ANSI
standard. Use of ANSI guidelines leads to maximum permissible exposure (MPE) values for
scanning microscopy that are unnecessarily low. The major damage mechanism for illumination
in the NIR wavelength region is known to be thermal. More accurate attempts to compute the
irradiance damage threshold for tissue thus consider the amount of energy deposited and the
temperature rise generated by a small focal volume. These models typically consider the
absorptive heating of specific chromophores in the tissue. To properly assess thermal damage,
however, the rapid cooling effects due to steep thermal gradients must also be considered.
Unfortunately, no comprehensive analysis yet exists for scanning microscopy. Determination of
exposure limits has instead been based to date on histologic analysis of irradiated tissues. In vivo
confocal microscopy systems imaging at 15 - 30 frames per second use up to 20 mW of incident
power at 1060 nm focused to a spot size of less than 1 micrometer [10]. At this power level,
human skin cells can be imaged over several minutes without damage. Assessment of exposure
limits for OCM imaging in this thesis were based on the published confocal work. Because the
111
OCM system currently operates at slower imaging speed than the in vivo confocal microscope,
the irradiance level is limited to below 10 mW as an added safety measure. A protocol for in
vivo imaging at this irradiance was approved by the MIT Committee on the Use of Humans as
Experimental Subjects (COUHES), and informed consent was obtained from volunteers before
commencing studies.
4.3.2
Tissue Stabilization
When imaging structure on the micron scale, motion artifact is a significant problem. High
imaging frame rate is used to combat some of the motion artifact on a single image basis, but
tracking of cellular features over the course of many images requires tissue stabilization
schemes. Stabilization for this thesis work consisted of immobilization of the specimen with a
simple clamping technique. Stabilization could be maintained reasonably well for single images,
but frame-to-frame tracking of features was inadequate. A more effective stabilization scheme
was demonstrated by Rajadhyaksha for in vivo confocal microscopy [10]. This stabilization
technique fixes the tissue in direct mechanical contact with the microscope housing. In addition,
it uses a coverslip to provide additional contact stabilization of the specimen. Future work will
include implementation of a contact stabilization scheme similar to that used for confocal
microscopy and reduction of single frame motion artifact through improvement of imaging
speed.
4.3.3
Imaging with Short Coherence Gate
Short coherence gate OCM imaging was investigated at 800 nm for various confocal gates to
determine if cellular resolution imaging can be obtained with a relaxed confocal gate. Use of a
small coherence gate to dominantly set the axial section thickness could alleviate the need for
very high numerical aperture optics in the sample arm, which would in turn enable several probe
designs that cannot be implemented at high NA.
Figure 4.4 provides images acquired with a 12 um confocal gate in combination with 3 urn
coherence gate. The images are of the nailfold region of a person of skin type IV. The setup
used the reflective microscope design at 800 nm together with the 30 X, 0.9 NA objective lens.
Incident sample power was 3 mW with measured sensitivity greater than 85 dB. Images were
acquired at 2 - 4 frames per second with a field of view of 145 um x 100 um. Image processing
consisted solely of adjustment of the aspect ratio and the gray scale.
Cellular features are clearly visible at varying layers of the epidermis. Figure 4.4a is near the
surface layer as evidenced by the large, puffy looking keratinocytes located either in the
cornified layer or in the upper granular layer. Cells captured in figures 4.4b - 4.4d are most
likely located in the lower granular layer or in the spinous layers, as evidenced by the slightly
smaller cell size. Figures 4.4e and 4.4f appear to be the papillary ridges between the dermis and
epidermis. The small cells on the ridges would then be the basal cells. These ridges are typically
located between 100 and 150 um below the surface. Slightly deeper, figures 4.4g and 4.4h show
capillary or duct structure in the superficial dermis. These observations correspond well to those
made with in vivo confocal microscopy.
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Figure 4.4. In vivo cellular images of human skin. Wavelength = 800 nm. Confocal gate =
12 um. Coherence gate = 3 um. Sample power = 3 mW. Image field of view is 145 um x
100 um. A - D acquired at 4 fps. E - H acquired at 2 fps.
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Note that the confocal section thickness of 12 um is not alone sufficient with regard to the 5
um requirement for imaging cellular structure deep in scattering tissue. The combination with a
3 um coherence gate, however, provides the additional required axial sectioning capability.
Furthermore, the transverse resolution is still sufficient to image cellular structure despite the
relaxed confocal parameter.
To investigate further the ability to image cellular structure with short coherence gate, the
OCM system was reconfigured with 30 um confocal gate and 3 urn coherence gate using the
close-coupled scanner microscope design. Figure 4.5 presents images obtained using this
configuration. The images are taken in the epidermis of the nailfold region of skin type III and
type IV volunteers. Incident sample power was around 5 mW leading to a sensitivity of
approximately 94 dB. Images were acquired at 4 frames per second with a field of view of 210
um x 144 um. Image processing consisted only of adjustment of the aspect ratio and
optimization of the gray scale.
Again cellular features are visible throughout the epidermis. Figures 4.5a through 4.5d
illustrate the granular and spinous cell layers while 4.5e and 4.5f show the ridges at the
epidermis-dermis junction. Figures 4.5g and 4.5h then show the presence of capillary or duct
structure again in the superficial dermis.
The 30 um confocal gate is prohibitively large for high contrast confocal imaging in highly
scattering media. This result clearly illustrates the enhanced sectioning provided by the short
coherence gate. The 3 um coherence gate sets the axial section thickness while the sample arm
optics merely are required to provide small spot size. Achievement of high transverse resolution
is a much less stringent optical design requirement than achievement of a small confocal gate.
4.3.4
Imaging Depth Measurement
The maximum imaging depth for the 800 nm OCM system was measured in vivo by
stabilizing the tissue and taking images at calibrated depths using a motorized stage. As was
mentioned in section 4.2, the method of tissue stabilization is not yet sufficient for reliable
tracking of features between frames. As such, a single continuous depth scan cannot be counted
as a reliable measure of the imaging depth. To combat against motion, the depth measurement
was made as a series of scans. An imaging depth was chosen, the stage translated to the desired
depth, and an image acquired. For the next desired depth, the focus was returned to the surface
for calibration and then translated to the correct depth again. The procedure was repeated until
the maximum imaging depth was determined. Figure 4.6 displays images taken deep in the
dermis of a person with type IV skin. Heterodyne signal is maintained until over 350 um below
the surface. This value is the actual stage translation. There is also a shift in the focus to greater
depths when imaging deep in tissue. This shift is caused by refraction of the converging beam at
the interface between the tissue and the immersion medium. Radjadhyaksha calculated this shift
for determination of imaging depth in human skin with in vivo confocal microscopy [10]. From
his calculation, a shift of the objective by 350 um corresponds to an actual imaging depth of 378
um. Note that the maximum depth of imaging at 800 nm using confocal microscopy is between
150 and 200 um [9]. OCM clearly provides capability for imaging deeper than standard confocal
microscopy.
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Figure 4.5. In vivo cellular images of human skin. Wavelength = 800 nm. Confocal gate =
30 um. Coherence gate = 3 um. Sample power = 5 mW. Image field of view is 210 um x
144 um. Images acquired at 4 frames per second.
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Figure 4.6. Imaging deep in the human dermis. Wavelength = 800 nm. Confocal gate = 30
um. Coherence gate = 3 um. Sample power = 5 mW. Image field of view is 210 um x 144
um. Images acquired at 4 frames per second.
It should be noted that the value of 378 um for imaging depth is likely conservative estimate
of the capability of the technology. During the measurement, no effort was made to ensure
coordination of the coherence and confocal gates as the objective was translated deeper into
shift. From the analysis of chapter 2, it is clear that the presence of refractive index mismatch
can cause the two gates to slip relative to each other and thereby degrade the heterodyne signal.
Furthermore, the motion artifact needs to be reduced via adequate tissue stabilization and higher
imaging speeds in order to more accurately measure imaging depth in vivo. Future work will
include design of control mechanisms for coordinating the overlap of the confocal and coherence
gates as well as implementation of improved stabilization schemes and higher imaging speed.
These advances will enable more precise measurements of imaging depth both in vitro and in
vivo.
4.4
Preliminary Imaging Results with a Handheld Probe at 1300 nm
The handheld imaging probe was integrated into the 1300 nm OCM system and
demonstrated for imaging of Xenopus laevis tadpole. The AFC superluminescent diode laser
source was used with 65 nm of bandwidth and a resulting coherence gate of approximately 12
um. The measured confocal gate for the probe was 15.5 um. Images were acquired at 4 frames
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per second with nearly 4 mW of power onto the sample. The preliminary results are shown in
figure 4.7.
Figure 4.7. In vivo imaging of Xenopus laevis using handheld probe at 1300 nm. Confocal
gate = 15.5 um. Coherence gate = 12 um. Images acquired at 4 frames per second with 4
mW power on sample. Field of view ~ 140 um x 140 um.
The high transverse resolution of the probe is evident in the images. Individual cell nuclei
and cell membranes are clearly distinguishable. The longer coherence and confocal gate
parameters, however, do not provide sufficient axial sectioning capability for imaging in highly
scattering tissues such as skin. Future work will involve incorporation of the broadband
Cr:Forsterite laser into the 1300 nm system together with the handheld probe. The Cr:Forsterite
laser can provide a coherence gate of around 5 um, which should be adequate for imaging in skin
and other human tissues. The longer wavelength will provide improved penetration and could
enable cellular imaging to depths approaching 1 mm, far greater than imaging depths achieved
with confocal microscopy alone.
117
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Chapter 5
Summary and Future Work
A novel, broadband optical coherence microscopy system has been demonstrated for realtime, in vivo cellular imaging in human tissue. The system uses a reflective grating phase
modulator in combination with femtosecond laser sources to achieve coherence gates of a few
micrometers. In this operating regime, the coherence gate provides sufficient axial sectioning for
imaging in highly scattering tissue even in the presence of relatively weak confocal sectioning.
The ability to relax the confocal gate translates to a lower numerical aperture requirement for
deep imaging of tissue microstructure. This result promises to enable cellular imaging with
miniaturized probe designs that are unsuitable for confocal microscopy yet essential for
important clinical applications such as endoscopy and catheterization.
The OCM system was demonstrated for imaging Xenopus laevis tadpole. High contrast
images of cellular detail were obtained. In addition, real-time acquisition was demonstrated by
imaging flow of blood cells in vessels of various sizes. These preliminary results emphasize the
high resolution capability of optical coherence microscopy and suggest a role in developmental
biology studies.
As a more rigorous test, real time OCM images were taken of human skin in vivo. Using a 3
um coherence gate, cellular structure was discernible throughout the epidermis with a confocal
gate of 30 um, much larger than the 5 um limit typically considered acceptable for in vivo
confocal microscopy. Sufficient transverse resolution for cellular imaging remained despite
weak confocal sectioning. Furthermore, imaging was demonstrated in the dermis to depths
beyond quoted values for confocal microscopy.
Future work will focus on continued optimization and characterization of system parameters
and transition of OCM technology to clinically relevant applications. For in vivo imaging, a
more effective tissue stabilization scheme based on those used for confocal microscopy will be
implemented. In addition, a mechanism for controlled coordination of reference and sample arm
path lengths will be developed. This is necessary step to combat the slip between confocal and
coherence gates occurring when focusing deep into inhomogeneous scattering tissue. These two
improvements will enable calibrated measurements of imaging depth in vivo, which will help in
quantifying advantages of OCM over confocal microscopy for imaging tissue microstructure.
Optimization of the electronic receiver will also improve system performance. Contrast and
resolution likely suffer in the current system with the use of a log demodulator. Since the
dynamic range of the OCM images should be smaller than that of OCT images, linear
demodulation with adjustable gain may be the better choice. Various approaches to linear
demodulation will be tested including analog quadrature demodulation and direct digitization of
the heterodyne interference signal followed by software demodulation.
Investigation of operating parameters for OCM will continue. In particular, further work will
seek to determine optimal coherence gate and confocal gate parameters for imaging in vivo.
Using a shorter confocal gate improves transverse resolution but also increases the sensitivity to
pathlength mismatch and the numerical aperture requirement on the probe. A longer confocal
gate, on the other hand, relaxes the optical design constraints on the microscope and reduces the
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path mismatch problem. This compromise will be further studied with respect to designated
applications of the technology. In addition, the dependence of OCM image quality on the
coherence gate will be studied by varying the light source bandwidth.
This thesis demonstrated in particular the advantages of the short coherence gate, long
confocal gate operating regime for OCM. Future work will also look at the long coherence gate,
short confocal gate regime, typically known as coherence-gated confocal microscopy. Imaging
depth and contrast will be measured and compared between the two working limits to more fully
characterize the advantages of the technology and its potential applications.
The broadband OCM system will be adapted for use with femtosecond lasers at longer
wavelengths, namely 1064 nm and 1300 nm. Longer wavelengths will provide increased
penetration and may enable imaging to depths approaching 1 mm or more. Confocal microscopy
systems do not typically use wavelengths beyond 1064 nm because the axial section thickness
Increasing wavelength only exacerbates the already stringent
scales with wavelength.
requirement for high numerical aperture. OCM should offer an advantage in this respect, since
section thicknesses approaching 5 um can be obtained with femtosecond sources independently
of the sample arm optics.
In vitro and in vivo imaging studies will continue on a variety of animal and human tissues.
In vitro OCM imaging of normal and pathologic tissue samples will help to define the diagnostic
potential of the technology. Based on these results, select in vivo clinical applications may be
chosen for further development. Animal studies will continue on developmental biology
specimens such as Xenopus laevis and Rana pipiens and will be extended to model systems for
disease development such as the hamster cheek pouch model for carcinogenesis.
Transition to in vivo clinical applications will require design of compact, robust probe
technology. Future work will include the development of a second generation handheld probe
and catheter based devices for endoscopic applications. Probe design will consider various
microscanning technologies, including piezoelectric elements and MEMS devices.
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