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27 - Photonic Materials, Devices and Systems – Laser Medicine and Medical Imaging Group – 27
RLE Progress Report 144
Laser Medicine and Medical Imaging Group
Academic and Research Staff
Prof. James G. Fujimoto
Visiting Scientists, Research Affiliates and Collaborators
Prof. Mark Brezinski, Prof. Joel Schuman, Dr. Wolfgang Drexler, Dr. Stephane Bourquin, Dr.
Christian Chudoba, Dr. Ingmar Hartl, Dr. Hiroshi Ishikawa, Dr. Christopher Kroeger, Dr. Xingde Li,
Dr. Adelina Paunescu, Dr. Ping Xue, Nirlep Patel, Karl Schneider
Graduate Students
Aaron Aguirre, Ravi Ghanta, Pei-Lin Hsiung, Paul Herz, Tony Ko, Costas Pitris
Technical and Support Staff
Cindy Kopf, Mary Aldridge, Donna Gale
Technology for Biomedical Imaging and Optical Coherence Tomography
Sponsors
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-5-R01-CA75289-04
National Institute of Health
Grant NIH-1-R01-AR44812-02
National Institute of Health
Grant NIH-1-R01-HL63953-01
National Institute of Health
Grant NIH-1-R01-EY131780-01
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Andrew M. Kowalevicz, Pei-Lin Hsiung, Paul Herz, Costas
Pitris, Dr. Stephane Bourquin, Dr. Christian Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Dr.
Hiroshi Ishikawa, Dr. Adelina Paunescu, Dr. Thomas Schibli, Dr. Ping Xue, Nirlep Patel, Karl
Schneider, Professor Franz X. Kaertner, Professor Mark Brezinski, Professor Joel Schuman, and
Professor James G. Fujimoto
Ultrahigh Resolution Optical Coherence Tomography
Optical coherence tomography (OCT) is an emerging technology for micron-scale cross-sectional
imaging of biological tissue [1, 2]. OCT is analogous to ultrasound imaging except that it uses
light instead of sound. Imaging is performed by measuring the echo time delay of light. Since the
speed of light is extremely rapid, it cannot be measured by direct electronic detection, but instead
relies on low coherence interferometer.
The longitudinal resolution in OCT images is inversely proportional to the optical bandwidth and
proportional to the square of the center wavelength of the light source. The axial resolution in low
2
coherence interferometry is related to the bandwidth of the light source according to λ /∆λ.
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Ultrahigh resolution OCT requires extremely broad bandwidths because of this λ /∆λ dependence
of the longitudinal resolution. This is particularly the case for the spectral region between 1.2 µm
and 1.5 µm. This spectral region is of high interest for OCT because of the high penetration depth
in biological tissue and the possibility to perform spectral resolved imaging of water absorption
bands. Superluminescent diodes are used in conventional OCT systems and typically yield 10-15
µm longitudinal resolutions. We demonstrated previously OCT imaging with resolutions of 1 µm at
800 nm [3] and 5.1 µm at 1300 nm [4] in biological tissue using a Kerr-lens modelocked
Ti:Sapphire laser with double-chirped mirrors and the self phase modulation broadened spectrum
of a Kerr-lens modelocked Cr:Forsterite laser, respectively.
2
These broadband Kerr-lens modelocked lasers are not available commercially and require
expensive and hard to fabricate double chirped mirrors. We tested a novel broad band light
source that is continuum generation in air-silica microstructured fiber for the application in OCT.
High nonlinearity, air-silica microstructure fibers [5] or tapered fibers [6] can generate an
extremely broadband continuum using low energy femtosecond pulses. These fibers achieve
high nonlinearities by using a tight mode confinement and shifting the zero of dispersion to
shorter wavelengths.
We successfully demonstrated the use of continuum generation in an air-silica microstructure
fiber for OCT imaging [7]. We achieved a free-space axial resolution of 2.5 µm at a 1.3 µm center
wavelength and demonstrated the application of the OCT system for in vivo imaging of biological
tissue. These are the highest resolutions achieved to date for imaging in this wavelength regime.
Figure 1 (left) shows a schematic of the experimental setup. We launched 100 fs pulses
generated by a commercial Kerr-lens modelocked Ti:Sapphire laser , which was pumped by a
frequency doubled Nd:Vanadate laser into a 1 m length of microstructure fiber. The fiber
continuum was collimated with a microscope objective (MO), spectrally filtered by a long-pass
filter and focussed with a chromatically corrected custom-designed lens (AL) into a dispersion
shifted single mode fiber. A custom-designed lens and broadband 3dB fiber couplers (FC),
designed for 1300 nm center wavelength (FC), were used in the OCT setup. Since the light
source had excess amplitude noise, a dual balanced detection with two InGAs photodiodes (D1,
D2) was used. Polarization controllers (PC) minimized the polarization mismatch of the
interferometer arms to avoid degradation of the shape and the peak height of the interference
fringes.
Figure 1. (left) Ultrahigh resolution OCT system using continuum generation in an air-silica
microstructure fiber as the light source. See text for explanation of the acronyms. (right) In vivo
ultrahigh resolution (~6 x 2.5 µm, transverse x longitudinal resolution, 1.0 x 1.0 mm; 1000 x 3000
pixels) OCT image of a Syrian hamster cheek pouch. Different layers are clearly visible.
We demonstrated the feasibility of in vivo ultra high resolution imaging by using the cheek pouch
of a Syrian hamster, a well-established animal model for studies of cancer progression. Figure 1
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(right) shows an in vivo image of cheek pouch which shows the epithelium, connective tissue and
muscle layers at ultrahigh resolution. The cheek pouch was index matched by using a
microscope cover glass and saline solution.
While this approach achieves high resolutions in the important 1.3 µm wavelength range, there
are several disadvantages. The femtosecond Ti:Sapphire laser used to generate the continuum
requires a high power pump laser which makes the light source expensive and bulky. In addition,
the microstructure fiber requires a free space coupling to the interferometer, which limits the
optical power at the sample and the stability of the system. Finally, in order to achieve broad
bandwidth, high-resolution operation, the OCT system used a delay line consisting of a
galvonometer scanned retroreflector. This limits the maximum axial scan speed and real time
imaging cannot be performed.
To this end, we have demonstrated a novel light source for OCT using a Nd:Glass femtosecond
laser with spectral broadening in a tapered single mode fiber. The Nd:Glass femtosecond laser is
much less expensive (~5 times less) than the Ti:Sapphire femtosecond laser and pump. In
addition, the Nd:Glass laser is compact (630mm x 320 mm x 160 mm) and can be more easily
integrated into a robust and portable system for OCT imaging in the clinical setting. We have
also designed a reflective grating phase delay line design that enables high speed, broad
bandwidth optical delay scanning with this system. High speed and high resolution OCT imaging
is achieved with 5 µm axial resolution and 90 dB sensitivity. Axial scan speeds of ~1 m/s are
demonstrated, sufficient for real time imaging.
OCT images taken with this light source and scanner were compared to images taken with a
conventional semiconductor light source. Figure 2 demonstrates this comparison on an
anesthetized African frog (Xenopus laevis) tadpole. The images consist of 676 axial pixels by
3000 transverse pixels, covering an area of 1 mm by 2 mm. The continuum light source yields an
axial resolution of 5 µm in air or ~4 µm in tissue, which is four times higher than the conventional
semiconductor source. This improved resolution enables the differentiation of small features
which are indistinguishable with conventional light sources.
Since the demonstration of OCT imaging with resolutions 5.1 µm at 1300 nm [4] in biological
tissue using the self phase modulation broadened spectrum of a Kerr-lens modelocked
Cr:Forsterite laser, we have worked toward the development of a compact and robust
Cr:Forsterite laser for future implementation in a clinical OCT system. Recently, we have
developed a compact (1’ x 2’ x 8”) Cr:Forsterite laser light source that can be integrated into a
clinical OCT system.
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0.25 mm
x4
0.25 mm
Figure 2. In vivo OCT image of an African frog tadpole (Xenopus laevis). (top) Conventional light
source (AFC technologies) with 20 µm axial resolution in air (16 µm in tissue). (bottom)
Continuum generated in tapered fiber as light source with 5 µm axial resolution in air (4 µm in
tissue).
We have also demonstrated ultrahigh resolution OCT imaging using fluorescence from a highly
doped Ti:Sapphire crystal. Using a 5 W cw laser pumping the crystal, 40.3 µW of fluorescence
with 138 nm bandwidth was coupled into a single-mode fiber (See Section I). This source
enables ultrahigh resolution OCT imaging with 2.2 µm axial resolution in air (1.7 µm in tissue) and
>86 dB sensitivity. Figure 3 demonstrates an OCT image taken with this source on an African
frog (Xenopus laevis) tadpole in vivo. The image is constructed using C-mode scanning of 5
OCT images taken at different focal planes. The image consists of 600 transverse and 625
longitudinal pixels, covering a field of view of 600 µm and 470 µm. The ultrahigh resolution OCT
image shows that individual cells and nuclei can be easily resolved. This demonstrates a simple
and robust light source for spectroscopy and ultrahigh resolution OCT imaging without the need
for complex femtosecond Ti:Sapphire lasers.
Figure 3. Ultrahigh resolution OCT image of an African frog tadpole (Xenopus laevis) taken with
Ti:Sapphire fluorescence. Imaging was performed with <2 µm axial resolution and 5 µm
transverse resolution. Five OCT images with different focal depth settings was fused to form this
image of extended depth of field
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Spectroscopic Optical Coherence Tomography
Spectroscopically resolved imaging is important because it promises to enable imaging of
functional properties of tissues on a micron scale by using spectroscopic properties. Relatively
few studies on spectroscopic OCT have been performed because broadband light sources have
not been available. An earlier study of spectroscopic OCT used a bandwidth of ~50 nm at 1.3
µm [8].
We recently used a Kerr-lens modelocked Ti:Sapphire laser with dispersion
compensation by double chirped mirrors for spectroscopic imaging in the wavelength range from
650 to 1000 nm [9]. Previous spectroscopic OCT studies in the spectral region of the water
absorption band at 1.45 µm used two different light sources in dual wavelength OCT systems.
One light source was used for absorption measurements and the second light source for
referencing [10, 11]. However, this approach adds the noise of two independent light sources,
and accurate referencing is hard to achieve.
In order to demonstrate spectroscopic OCT imaging, we investigated differential water absorption
by selecting a spectral range of 200 nm centered at 1400 nm from the continuum. We imaged a
phantom consisting of two microscope cover glasses of 170 µm thickness confining a water
sample. A schematic of the setup is shown in Figure 4 (left). The cover glasses were in contact at
one side of the sample and had a spacing of ~0.8 mm at a lateral distance of 10 mm. The focal
position of the imaging beam was ~0.4 mm above the bottom coverglass.
Figure 4. (left) The phantom consists of two microscope cover glasses enclosing H2O as
absorber. (center) Digitally demodulated OCT image of the phantom used for this study. (right)
Spectrum of the OCT echo of the cover glass surface at the points indicated with A and B and
one intermediate point. The arrow indicates wavelength of the water absorption band.
The full interference signal was digitally recorded at 12 bit resolution with 30,000 digitalization
points per 1.3 mm A-scan. A digitally demodulated image of the sample is shown in Figure 4
(center). The spectrum of several A-scans was calculated by Fourier-transforming a 2000 point
window centered at the maximum of the reflection signal. The spectra obtained were normalized
to the intensity maximum. The spectra at the points A and B and one intermediate point are
shown in Figure 4 (right). One clearly observes that the absorption of the spectral intensity
around 1.45 µm increases with increasing absorber thickness.
Novel Imaging Devices for Optical Coherence Tomography
The success of optical coherence tomography in clinical applications will depend in large part on
the design and availability of delivery mechanisms that allow seamless integration of OCT with
existing diagnostic imaging modalities. OCT has the advantage that it is fiber optically based and
can readily be integrated with a wide range of existing clinical imaging instruments such as
microscopes, laparoscopes, endoscopes, and catheters. As an imaging technique capable of
imaging human tissue at high resolution, OCT could lead to detection of pathology at earlier
stages than currently possible, leading to improved patient prognosis. Several imaging devices
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have been designed and evaluated to investigate the feasibility of OCT for the clinical
assessment of pathology of a variety of non-transparent tissue in vivo.
Linear and rotational scanning catheters
Since optical coherence tomography is fiber optically based, it can be easily integrated into a
wide variety of clinical imaging devices. We have developed a series of OCT catheters that are
being used in several ongoing clinical studies. We have developed a small diameter OCT
catheter, which can be used either for linear scan or radial scan. The OCT catheter consists of
an optical coupling element at its proximal end, a single-mode optical fiber running its length, and
optical focusing and beam directing elements at the distal end. The single-mode optical fiber is in
a flexible hollow metal cable that is either rotated or translated with a motor drive unit at the
proximal end. A schematic and a photograph of the distal end of the OCT catheters is shown in
Figure 1. The distal end contains a gradient index lens and a right-angle microprism to focus and
direct the beam at a 90° angle with respect to the axis of the catheter. For our studies we chose
the optics for the catheter to achieve a spot size (transverse resolution) of ~13 µm and a ~2-3 mm
working distance. The fiber optic cable and distal optics are protected by a sealed transparent
plastic sleeve which remains stationary as the fiber optic cable and optics move inside. The
catheter has an outside diameter of ~1 mm and can be passed through the working channel of a
standard endoscope. The linear scan probe scans the longitudinal position of the OCT beam at a
fixed angle, to generate a rectangular image of a longitudinal plane at a given angle with respect
to the probe. The radial scan probe directs the OCT beam radially outward and scans the angle
of emission rotationally to generate an image which is displayed in a polar plot.
transparent
guard microprism
SM fiber
sheath
500 µm
Guard
......
hollow
metal cable
Photo of Distal End
(on US Penny)
Aiming Beam
GRIN lens
silicone
sealant
Figure 5. Schematic and photograph of the OCT catheter distal end. The metal speedometer
cable, focusing gradient index lens, and right-angle prism are protected by a metal guard and
encased within a transparent sheath.
Both linear and radial-scanning OCT catheters have their advantages and limitations. In OCT
imaging catheters, the focused beam spot size (transverse resolution) and working distance are
determined by the distal micro-optics. The beam is in focus when the tissue is positioned near
the focal position of the beam within the depth of field. Images acquired with the linear-scanning
catheter are displayed in Cartesian (rectangular) coordinates, whereas images acquired with the
radial-scanning catheter are displayed in polar (circular) coordinates. The linear scanning
approach has the advantage that the pixel spacing in the transverse direction is always uniform.
This contrasts with the radial imaging approach where the transverse (circumferential) pixel
spacing increases with increasing distance from the catheter. Thus, the radial OCT images
become progressively coarser when large diameter lumens are scanned. To image with the
linear scan catheter design, the distal end is positioned against the wall of the lumen. Placing the
catheter in contact with tissue not only fixed the distance to the tissue, but also helped stabilize
the catheter position during cardiac or respiratory movements.
The radial scan method is analogous to that used in intravascular ultrasound catheters (IVUS)
which contain a rotating distal ultrasonic transducer. The radial scanning catheter is well suited
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for imaging small (few mm) diameter lumens such as human arteries or the gastrointestinal tract
of animals. However for large diameter lumens such as the human esophagus, it is difficult to
position and stabilize the radial-imaging catheter in the center of the lumen. Movements of the
lumen or endoscope resulted in motion artifacts and variations in the focus position. If radial
imaging were performed with the catheter de-centered in the esophagus and near the mucosal
surface, only a small sector of the 360° scan range would be in focus.
Doppler imaging catheter
It is possible to combine functional as well as structural imaging. For intravascular imaging, the
measurement of blood flow is especially important for assessing athersclerotic narrorwing of
arteries and the effects of atheractomy therapy. We have developed a Doppler OCT catheter
integrates functional imaging with structural imaging using OCT. The catheter consists of an
angle-cleaved single mode optical fiber with a gradient index (GRIN) lens glued to the angle
cleaved end of the fiber to focus the beam. In order to detect a Doppler signal due to flow, a nono
zero projection of the imaging beam along the longitudinal axis of the catheter is needed. A ~60
o
microprism is custom-made and employed so that the beam exits the catheter at ~60 with
respect to the catheter axis. The single mode fiber, the GRIN lens and the microprism are
attached to form a single unit and housed concentrically in a 27 Gauge (~410 µm) hypodermic
metal guard. The entire catheter is housed within a transparent plastic sheath with a sealed distal
end. At the proximal end of the catheter the hollow wire is connected to a DC motor. The optical
beam can be scanned radially by rotating the hollow metal wire along with the single mode fiber
and the distal optics. A rotary coupler based upon a glass capillary tube is used to connect the
rotating fiber at the proximal end of the catheter and a stationary fiber from the light source [12].
Figure 6 shows one representative schematic of the catheter and a schematic of the catheter
demonstrated for measuring the intraluminal flow profile in phantom.
Catheter
Sheath
beam
60o
3.17 mm
0.41 mm
1.14 mm
A
B
flow
tube
Figure 6. (a) Schematic of the Doppler OCT imaging catheter. The beam exits from the catheter
o
o
at an angle of 48 in air (60 in water) with respect to the catheter’s longitudinal axis. (b)
Schematic of the experimental phantom consisting of a clear plastic tube with an inner diameter
of 3.17 mm. 1% Intralipid is circulating through the tube at a constant flow rate.
The Doppler catheter can be used for assessment of intraluminal blood flow within a vessel as
well as structural imaging of the fine structures within the vessel wall. The Doppler OCT catheter
can also be potentially used for evaluating the severity of stenosis and the outcome of
cardiovascular interventions such as stenting. We are currently evaluating the potential of this
device for deducing the vessel wall diameter in a physiological relevant environment (blood).
This device can also be used for imaging the fine structures within the vessel wall when moderate
saline flushing is used. In addition to the flow velocity profile, the flow turbulence may also be
quantifiable by investigating the flow velocity variance. Future work will also focus on
investigating the potential of this device for diagnosis of thrombosis or aneurysm from the velocity
variance assessment.
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OCT Biopsy Needle
OCT applications in the past have been limited to the surfaces or lumens of organ systems
because of the ~2-3 mm penetration depth of OCT in most tissues. Pathologies within solid
tissues or organs have been outside the range of OCT. There are, however, many clinical
scenarios where high resolution imaging of solid tissues is desirable. One promising application
of OCT is in imaging pathology and guiding biopsy in solid tissues. This procedure could reduce
the sampling error of excisional biopsy in diagnosing cancers in solid organs such as the prostate
or breast. Other applications include optical imaging in scenarios where excisional biopsy is
hazardous, and in surgical guidance such as in cryosurgery or interstitial photodynamic therapy.
We have developed a prototype of an interstitial imaging needle for OCT that has a diameter as
small as 27 Gauge (~410 µm) [13]. A GRIN lens and a micoprism are used to focus and deflect
the optical beam which is emitted at an angle of 90 degrees from the axis of the needle. The
optical fiber is housed concentrically within a thin wall stainless steel hypodermic tube with a 27
gauge outer diameter. This hypodermic tube protects the optical fiber and facilitates the insertion
of the needle into the tissue. The imaging needle has a confocal parameter of ~380 µm,
corresponding to a spot size (or transverse resolution) of ~17 µm, with a focal distance ~80 µm
outside the optical window. The optical beam can be scanned radially by rotating the entire
needle along with the distal optics. This technique is similar to that used with acupuncture
needles. The OCT imaging plane is perpendicular to the needle axis and the position of the
imaging plane can be controlled by varying the depth of needle insertion.
The OCT imaging needles can be inserted into tissue independently or coupled to standard
biopsy techniques, such as fine needle aspiration (FNA) or core biopsy. Figure 7 shows one
concept of OCT guided biopsy. The OCT needle is introduced first, prior to the biopsy device. If
an area of tissue pathology is detected, the biopsy device would then be inserted in parallel,
riding the OCT needle in a monorail-like fashion. This approach has the advantage of allowing
selective insertion of the larger 18 gauge biopsy devices only when deemed necessary by the
imaging. Since the small 27 gauge OCT imaging needle would produce only minimal trauma,
more needle insertions could be performed than currently possible, allowing larger volumes of
tissue to be screened.
Biopsy
Imaging
biopsy
specimen
Imaging
Volume
Figure 7. Monorail biopsy and OCT imaging needle. A monorail design takes advantage of the
small size (27 gauge) of the OCT imaging needle, allowing imaging to be performed prior to the
introduction of an 18 gauge biopsy needle. This approach reduces trauma, allowing multiple sites
to be imaged before a decision to biopsy is made.
Needles of even smaller diameter (200 µm or less) can ultimately be built. The movement of the
needle can be performed with minimal resistance from the tissue and minimal trauma. OCT
imaging can assess larger volumes of tissue than are possible to excise and imaging can be
performed in real time prior to tissue biopsy with minimal tissue trauma. The OCT imaging
needle and OCT needle biopsy device would permit the imaging of pathology inside virtually any
solid organ or tumor and permit the guidance of biopsy based on microstructural features. This
would be an enabling technology for a wide range of research and clinical applications.
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Surgical Imaging Probe
The development of intraoperative OCT imaging enables pathology to be visualized in situ and in
real time in tissues that were previously inaccessible in vivo. Techniques that can visualize
microscopic disease in situ and in real time promise to improve diagnostic accuracy, and the
ability to visualize tissue structures intraoperatively should also facilitate the guidance of
microsurgical procedures. Imaging in the surgical suite requires a delivery device that is compact,
robust, and sterilizable using standard hospital sterilization techniques.
We have designed and developed a handheld probe and demonstrated its application for imaging
the prostate during nerve-sparing radical prostatectomy and for imaging cartilage during open
field surgery. The compact hand held probe consists of a Hopkins lens relay and a galvanometer
mirror which steers the beam in the transverse direction with the scanning range and spot size
magnified by telescope optics. The current imaging probe has a 2 cm working distance which
allows the surgeon to maneuver the beam within the tight confines of the surgical field. A low
power green aiming beam allows the surgeon to localize the imaging area and facilitate marking
for correlation with histology. The probe has a stainless steel outer tube with a transparent
window which can be sterilized using standard hospital sterilization procedures and slipped over
the distal end of the probe prior to imaging. The proximal end of the probe including cables can
be enclosed in a sterile surgical camera bag. The current probe scans over 3.5 mm and
produces a focused spot with an 18 µm transverse resolution. The transverse resolution and
scan range are determined by the choice of lenses in the relay system. If necessary, a small
coherent fiber bundle fiberscope can be used in conjunction with the hand held probe to allow
direct visualization of the region being OCT imaged.
A version of the probe has recently been designed that uses two orthogonally positioned
galvanometer mirrors to enable scanning of the OCT beam in the en face plane. Current studies
are underway using this probe for in vivo investigations of optical coherence microscopy, a
technique that combines standard OCT imaging with confocal microscopy. An additional ongoing
study includes using the probe with a portable OCT system at the National Cancer Institute to
image oral leukoplakia and monitor pharmacological chemoprevention therapy.
Optical Coherence Microscopy
Sponsors
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-5-R01-CA75289-04
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Pei-Lin Hsuing, Paul Herz, Dr. Stephane Bourquin, Dr.
Christian Chudoba, Dr. Ingmar Hartl, Dr. Ping Xue, Nirlep Patel, Karl Schneider, Professor Mark
Brezinski, and Professor James G. Fujimoto
Optical Coherence Tomography (OCT) has exciting potential for high resolution, in vitro and in
vivo optical biopsy in several types of scattering tissue [2]. OCT combines the coherence gating
afforded by broadband light sources with heterodyne detection to provide superior signal to noise
and image contrast at a range of wavelengths. Standard clinical OCT provides cross-sectional
imaging with 10-15 µm axial and transverse resolution. High-resolution OCT has 1-2 µm axial
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resolution and 5-6 µm transverse resolution, but is limited in the axial scan range due to the small
confocal parameter of the sample beam. Like OCT, laser scanning confocal microscopy (LSCM)
has also been demonstrated as a potential technique for high-resolution imaging of tissue
microstructure [14]. LSCM uses high-magnification, high numerical aperture objective lenses to
provide 1-3 µm transverse resolution and better than 1 µm axial sectioning capability. In contrast
to OCT, confocal microscopy samples an en face scan plane. Figure 8 compares cross-sectional
and en face imaging planes.
Transverse
X
Transverse
X
DEPTH PRIORITY
EN FACE
Incident
Beam
Incident
Beam
z Axial
z Axial
(Range Depth)
y
(Range Depth)
y
Figure 8. Schematic comparing scanning modalities for Optical Coherence Tomography (OCT)
and Laser Scanning Confocal Microscopy (LSCM). OCT uses depth scanning to form crosssectional imags. LSCM uses transverse scanning to form en face images.
LSCM has shown particular promise for in vivo imaging of cellular structure in a human tissues
[14], but signal attenuation in scattering media limits imaging depths to less than 1 mm. To
extend and improve upon confocal microscopy, our group demonstrated in 1994 the concept of a
new technique called optical coherence microscopy (OCM) [15]. OCM combines the coherence
gate with the confocal gate to improve rejection of unwanted scattered light that degrades image
contrast. With the superior signal to noise ratio offered by coherence-gated techniques, OCM
can yield approximately 2x greater imaging depths than standard confocal microscopy. Realtime, in vivo OCM has since been demonstrated for imaging cellular structure in human skin [16].
As with axial resolution in OCT, rejection of unwanted scattered light in OCM improves with
shorter coherence length, higher bandwidth light sources. Our recent work on high-resolution
OCM has focused on extending the technique to perform high speed imaging with broadband
light sources at multiple wavelengths. One major current limitation to achieving high resolution
OCM over a range of wavelengths is the lack of high speed, broadband phase modulators.
Previous studies have used either stretched fiber phase modulators, which have limited scanning
speeds, or waveguide phase modulators, which have limited bandwidths. Our group developed
and demonstrated grating phase delay scanners for high speed OCT imaging [17]. These
systems are similar to grating-based pulse amplitude and phase shaping techniques in
femtosecond optics [18]. The devices use a grating and lens to produce spectral dispersion
combined with a galvonometer-controlled mirror to produce a phase shift in the Fourier plane,
resulting in a group or phase delay. The grating phase delay scanner enables group and phase
delays to be independently controlled. Recently, grating phase delay scanners have been
demonstrated to achieve phase modulation [19].
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T o p V iew
C urved m irro r
G ratin g
G ratin g P h ase M o d u lato r
In R eflective G eo m etry
D ou b le
p ass
m irro r
B ro ad b an d S o u rce
+-
S can n in g
M irro r
D etection
E lectro nics
D1
In p ut b eam reflected
fro m m irro r
D o u b le
p ass
m irror
50/50
D2
M o n ito r
S id e V iew
C u rved m irro r
P o larizatio n C on trol
C o m p uter
M irror
S can n in g
M irro r
S -V H S
R eco rd er
f2
f
f3
G ratin g
f1
f2 + f3
4
F u n ctio n
G enerato r
G alvo
C on trollers
x
y
m o d u lato r
sam ple
O b jective
30x 0.9 N A
f3
H an d -H eld X -Y S can n in g P ro b e
f2
y-g alvo
x-g alvo
Figure 9. Optical coherence microscopy system using a reflective phase delay scanner and a
compact hand-held probe. A broadband light source is split into reference and sample beams by
a 50/50 fiber optic coupler. The sample beam passes through a reflective grating phase
modulator while the sample beam is focused onto the specimen using a novel hand-held probe.
The returned beams are recombined at a dual-balanced electronic receiver integrated with a PC
and S-VHS recorder. The reflective grating phase modulator supports broad bandwidth, high
axial resolution imaging and also enables real time imaging at a wide range of wavelengths.
We designed and demonstrated OCM using a novel reflective grating phase modulator and a
compact handheld probe. Figure 9 shows the experimental setup for the optical coherence
microscopy system. Our grating phase modulator is similar to reflective pulse stretchers and
compressors previously developed for femtosecond pulse amplification [20]. The reflective phase
delay line can support broad spectral bandwidths and perform high-speed phase modulation. This
system design will enable high resolution, high speed OCM imaging at wavelengths which were
previously inaccessible.
We acquired preliminary data using a superluminescent diode laser source with center
wavelength of 1300nm and 65 nm bandwidth. We achieved transverse resolution of better than
3.3 µm and demonstrated in vivo imaging on an African frog tadpole (Xenopus laevis) (Figure
10). The tadpole was imaged from the dorsal side with 1375x500 and 1375x250 pixel resolutions
at 2 to 4 frames per second, respectively. Cellular structure was clearly visible at multiple en face
imaging depths. The cell nuclei and cell borders appear highly scattering. In addition, circulatory
flow in a large vessel was also visible in sequential images or video. To demonstrate materials
imaging, OCM imaging was performed on laser fabricated optical waveguides in glass. The
waveguides are inside the glass and normally require phase contrast microscopy for visualization.
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Figure 10. OCM en face images taken of Xenopus laevis tadpole and of a laser-fabricated
waveguide embedded in glass (c). Images a, d-e are 1375x250 pixels taken at 4 frames per
second, while images b and c are 1375x500 pixels taken at 2 frames per second. The field of
view in the images is 130 µm x 140 µm.
We will combine our work on broad-bandwidth light sources with our new system design to
achieve high-resolution OCM at multiple wavelengths. We believe the system will provide a
powerful tool for optical biopsy to explore the limits of coherence-gated imaging for clinical
diagnostic purposes.
Ultrahigh Resolution Ophthalmic Imaging
Sponsors
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-1-R01-EY131780-01
National Institute of Health
Grant NIH-5-R01-CA75289-04
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
Air Force Office of Scientific Research
Grant F49620-98-01-0084
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Andrew M. Kowalevicz, Pei-Lin Hsuing, Paul Herz, Dr.
Stephane Bourquin, Dr. Christian Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Dr. Hiroshi
Ishikawa, Dr. Adelina Paunescu, Dr. Thomas Schibli, Dr. Ping Xue, Nirlep Patel, Karl Schneider,
Dr. Uwe Morgner, Professor Franz X. Kartner, Professor Mark Brezinski, Professor Joel Schuman
(New England Eye Center), and Professor James G. Fujimoto
OCT has perhaps been most widely investigated in ophthalmology, where it is beginning to make
a clinical impact in the assessment of retinal diseases such as macular holes, age-related
macular degeneration, glaucoma, and diabetic retinopathy [1, 21-23]. Current clinical practice
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calls for the development of techniques to diagnose ophthalmic disease in its early stages, when
treatment is most effective and significant irreversible damage can either be prevented or
delayed. With typical axial resolutions of 10 µm, OCT already provides more detailed structural
information than any other conventional imaging technique. However, the detection of many of
the early changes associated with diseases can require more accurate quantification of retinal
structure than is possible with standard resolution OCT.
250
µm
Figure 11. In vivo standard resolution (top) and ultrahigh resolution (bottom) OCT images of a
normal human fovea at approximately the same site. Resolutions were 10-15µm (axial) x 15µm
(transverse) and 3µm (axial) x 15µm (transverse) respectively.
Using the broad bandwidth of our ultrahigh resolution OCT system, we can image with axial
resolutions greater than 3 µm (in the retina), corresponding to a factor of 5 improvement over
OCT technology using superluminescent diode sources. The signal to noise ratio for the system
is ~100 dB. This system enables a significant improvement in the visualization of intraretinal
structures for earlier diagnosis and more precise staging of pathology (Figure 11). To our
knowledge, the image shown in Figure 11 represents the highest resolution in vivo image ever
acquired of the human retina. Ultrahigh resolution OCT offers an unprecedented visualization of
retinal morphology with structures such as the retinal nerve fiber layer, retinal pigment epithelium,
and the inner and outer plexiform layers. These structures are relevant in a variety of retinal
diseases, including age-related macular degeneration, diabetic retinopathy, and glaucoma (the
three leading causes of blindness worldwide).
Nasal
Temporal
Thickness (µm)
250
µm
350
250
Henle’s and Photoreceptor Layer
Thickness
Retinal Thickness
150
50
-3.0
Ganglion/NFL Layer Thickness
0
3.0
Transverse Position (mm)
Figure 12. Ultrahigh resolution allows for an unprecedented visualization of intraretinal structures
that may be quantified to provide an objective measure of retinal disease.
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Image processing techniques can be applied to acquired tomograms to quantify retinal and
intraretinal structures relevant to disease. Figure 12 illustrates the application of preliminary
image processing segmentation algorithms to ultrahigh resolution OCT images to quantify retinal
and intraretinal structures. Ultrahigh resolution enables the quantification of layers relevant to
retinal disease, which were previously not visualized or quantified using standard resolution OCT.
Precise quantification of the retinal thickness is important for the diagnosis and staging of macular
edema and diabetic retinopathy. Quantification of the photoreceptor and Henle’s layer may be
important in a variety of retinal diseases. Quantification of the ganglion cell layer and the nerve
fiber layer is important in retinal diseases such as glaucoma. Figures 13-15 demonstrate the
ability use the ultrahigh resolution OCT to quantify and map the different retinal layers in a normal
volunteer.
1.5
Superior
0.5
Retinal
Thickness
(µm)
Nasal
0
Temporal
Horizontal Position (mm)
1.0
-0.5
-1.0
-1.5
-1.5
Inferior
-1.0
-0.5
0
0.5
1.0
1.5
Transverse Position (mm)
Figure 13. In vivo two-dimensional retinal thickness map of a normal human macula. This image
is constructed from 4200 measurement locations (7 tomograms consisting of 600 A-Scans each).
The transverse sampling resolution was 5 µm and the horizontal sampling resolution was 500 µm
per pixel.
Superior
1.0
0.5
0
Nasal
Center
Horizontal Position (mm)
1.5
-0.5
Nerve Fiber
Layer
Thickness
(µm)
-1.0
Inferior
-1.5
0
1.0
2.0
3.0
4.0
5.0
Transverse Position (mm)
Figure 14. In vivo two-dimensional nerve fiber layer thickness map along the papillomacular axis
of a normal human volunteer. This image is constructed from 6000 measurement locations (10
tomograms consisting of 600 A-Scans each). The transverse sampling resolution was 5 µm and
the horizontal sampling resolution was 520 µm per pixel.
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1.5
Superior
0.5
Nasal
0
Temporal
Horizontal Position (mm)
1.0
GCL+IPL
Thickness
(µm)
-0.5
-1.0
-1.5
-1.5
Inferior
-1.0
-0.5
0
0.5
1.0
1.5
Transverse Position (mm)
Figure 15. In vivo two-dimensional intra-retinal layer thickness map of a normal human macula.
This image in constructed from 4200 measurement locations (7 tomograms consisting of 600 AScans each). The transverse sampling resolution was 5 µm and the horizontal sampling
resolution was 500 µm per pixel.
The visualization and quantification of retinal and intraretinal layers should serve as a valuable
clinical tool for the early assessment of ophthalmic disease. This concept has already been
demonstrated in a mouse retinal disease model, which allows us to follow and track different
retinal diseases in this animal. Using the ultrahigh resolution OCT system, we have imaged and
identified the many intraretinal layers of the mouse retina. Figure 16 illustrates the ability for
ultrahigh resolution OCT to visualize the intraretinal layers of a normal mouse retina in vivo.
When compared with histology taken from the same animal, the ultrahigh resolution OCT image
corresponds well with the layers identified in the histology.
NFL
IPL
INL
OPL
ONL
RPE
100 µm
Figure 16. In vivo ultrahigh resolution OCT image of a normal mouse retina and corresponding
histology. The layers identified in the OCT images correspond well with the layers in the histology
In the rhodopsin knockout mouse, the outer plexiform and outer nuclear layers undergo
degeneration three months postpartum. Figure 17 illustrates the differences between the in vivo
OCT images of a normal mouse retina and a rhodopsin knockout mouse retina. At 5 months of
age, the outer plexiform layer and the outer nuclear layer of the rhodopsin knockout mouse retina
would have undergone degeneration. When comparing the knockout mouse retina with the
normal wild type mouse retina, the OCT image clearly demonstrates this degeneration in the
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knockout mouse. The in vivo ultrahigh resolution OCT images clearly depict the degeneration of
the outer plexiform and the outer nuclear layer in the rhodopsin knockout mouse retina.
IPL
INL
OPL
50 µm
ONL
RPE
Rd -/- Knockout
(5 months)
Rd +/+ W ildtype
Figure 17. In vivo ultrahigh resolution OCT image of a normal mouse retina (Rd +/+ wild type) and
a rhodopsin knockout mouse retina (Rd -/-). OCT has the ability to quantify the thickness of the
different intraretinal layers as well as track disease progression in a non-invasive manner.
In addition to in vivo imaging of retinal in animal models and patients, we have also investigated
the ability of ultrahigh resolution OCT to discriminate the microscopic anatomy of the human
retina in vitro, and we compared ultrahigh resolution OCT measurements to the
histomorphometry taken from the same cadaver eye. Both normal and glaucomatous cadaver
eyes were obtained from National Disease Research Interchange (NDRI) and New England
Medical Center (NEMC) were fixed in a 3% glutaraldehyde solution. To minimize subretinal fluid
after sectioning the eye, perfluorocarbon solution was used to flatten the retina. The retinal
thickness was measured directly on the OCT scans taken within 48 hours of death. The
corresponding histomorphometry was performed on this histology using a Nikon Elipse E400
microscope. Figure 18 shows comparison between the ultrahigh resolution OCT image and the
corresponding histopathological image obtained with the light microscope on a glaucomatous
human cadaver eye.
Figure 18. Comparison between ultrahigh resolution OCT image and histomorphometry on optic
disk of a glaucomatous human cadaver eye.
Retinal layers, including the nerve fiber layer, inner nuclear layer, outer nuclear layer, the retinal
pigment epithelium, choriocapillaris and choroids could be identified. The correspondence of
retinal layers between the ultrahigh resolution OCT image and the histopathology is presented in
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Figure 19. Ultrahigh resolution OCT and histomorphometry produced similar retinal thickness
data in normal and glaucomatous eyes. We found differences in retinal thickness between
normal and glaucomatous eyes in both ultrahigh resolution OCT and histomorphometry.
Figure 19. Correlation between retina layers obtained with the ultrahigh resolution OCT and light
microscopy on glaucomatous human cadaver eye. The images represent the same section of the
retina of the same eye.
These studies demonstrate that ultrahigh resolution OCT can be used to identify different
intraretinal layers, quantify layer thickness, and track retinal disease progression. Future work
will involve ultrahigh resolution OCT imaging studies of patients with retinal diseases such as
age-related macular degeneration, diabetic retinopathy, and glaucoma.
Ultrahigh-Resolution and Spectroscopic OCT Imaging of Gastrointestinal
Malignancies
Sponsors
National Institute of Health
Grant NIH-5-R01-CA75289-04
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
Air Force Office of Scientific Research
Grant F49620-98-01-0084
National Institute of Health
Grant NIH-2-R01 EY11289-15
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Andrew M. Kowalevicz, Pei-Lin Hsiung, Paul Herz, Dr.
Stephane Bourquin, Dr. Christian Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Dr. Ping Xue,
Nirlep Patel, Karl Schneider, Dr. Hiroshi Mashimo and Dr. Muthoka Mutinga (VA Medical Center),
Dr. Jacques Van Dam (Stanford U.), Professor Mark Brezinski, and Professor James G. Fujimoto
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Endoscopic OCT has already been performed by our group using linear and radial scanning
methods to demonstrate the ability to discern changes in architectural morphology associated
with Barrett’s esophagus [24]. This study has demonstrated that endoscopic OCT can
differentiate normal from Barrett’s epithelium in real-time based on differences in epithelial
architecture. This suggests that OCT may be used as an adjunct to endoscopy for screening and
surveillance of Barrett’s esophagus, in particular in situations where endoscopic visualization is
difficult, such as to detect short segment Barrett’s or as a follow up to photodynamic therapy to
detect residual islands of Barrett’s pathology. However, the image resolution of conventional OCT
(~10-15 µm) is sufficient to differentiate architectural but not cellular morphology.
We have performed in vitro ultrahigh-resolution OCT and spectroscopic OCT on human
esophagus as a first step towards an ultrahigh resolution and spectroscopic endoscopic OCT
imaging clinical system. Ultrahigh resolution potentially allows visualization of tissue structure at
the cellular level, which could facilitate Barrett’s detection. Spectroscopic OCT was investigated
as a modality for improving contrast and differentiation OCT images. This technique may help
differentiate regions of different spectroscopic optical backscattering properties that are
associated with different tissue morphologies.
Imaging of in vitro specimens was performed using a bench top ultrahigh-resolution OCT system
[3]. This state-of-the-art OCT system performs imaging using a femtosecond Ti:Al2O3 laser light
source that emits ~5 fs optical pulses, corresponding to bandwidths of up to 350 nm at a center
wavelength of 800 nm. The OCT system was designed to support up to 260 nm of bandwidth,
achieving 1.5 µm axial resolution in free-space or ~1.1 µm in tissue. This is an order of
magnitude improvement in resolution compared to conventional OCT systems. An achromatic
lens with a 10 mm focal length was used to focus the beam on the in vitro tissue specimens to
achieve 5 µm transverse resolution. This system was used to image in vitro Barrett’s esophagus
specimens to demonstrate the potential of ultrahigh-resolution OCT imaging.
Spectroscopic OCT imaging is an extension of ultrahigh-resolution OCT which utilizes the broad
spectral bandwidth of the optical source to obtain information from the spectral content of the
backscattered light [9]. Conventional OCT imaging detects the envelope of the interference signal
which is generated by the reflections of light from the reference and sample arms. The amplitude
of this envelope is used to display the intensity of optical backscatter from specific regions within
the tissue. Spectral information can be obtained by digitizing the full interference signal and using
digital signal processing techniques.
Three surgical specimens of human esophagus were used. Imaging was performed only on
discarded tissue specimens. Specimens were maintained in a 0.9% saline solution and imaged
with ultrahigh-resolution and spectroscopic OCT within 3 hours of resection. The tissue specimen
was placed on a three-dimensional, computer-controlled, micron-precision translation stage.
Imaging was performed using the OCT microscope beam delivery apparatus. Immediately
following OCT imaging, the specimen was marked with India ink spots to indicate the OCT image
plane location. Specimens were placed in 10% Formalin for at least 24 hours prior to standard
histological processing. Histological slides were prepared from 5 µm thick sections stained with
hematoxylin and eosin. OCT images were compared with corresponding histological images.
Figure 20 (a) is an OCT image of Barrett’s epithelium. The image size was 2 mm x 1 mm (1000 x
1500 pixels). The image resolution was 1.1 µm axial and 5 µm transverse. Figure 20 (b) shows
corresponding histology from the same site. Multiple crypts and glandular structures were
observed along the surface. Magnified images of selected regions are shown in Figure 20 (c)
and 20 (d). Isolated regions of low and high optical backscatter were observed. Regions of low
backscatter corresponded to fluid-filled crypts or glands. It is unclear if individual cellular features
are visible, although the localized points of high backscatter are suggestive of nuclei.
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D
C
A
300 µm
C
100 µm
B
300 µm
D
100 µm
Figure 20. Ultrahigh-resolution OCT imaging of Barrett’s esophagus in vitro. (a) OCT image
acquired at ~1.1 µm axial and 5 µm transverse resolution illustrating multiple crypt- and gland-like
structures. (b) Corresponding histology confirms architectural morphology observed in the OCT
image. (c,d) Magnified regions from sites indicated in (a). Localized regions of high optical
backscatter (dark pixel clusters in image) may represent individual cell nuclei.
Spectroscopic OCT was performed on regions of normal and Barrett’s esophagus in vitro to
assess if spectroscopic imaging could enhance image contrast or improve the differentiation of
pathology. Amplitude tomograms represent the intensity of backscatter light, while spectroscopic
tomograms (Figure 25) measure the spectral content of the backscattered light. For the
spectroscopic OCT images presented here, the spectral shift of the backscattered light was
measured by calculating the center of gravity of the spectrum from each pixel in the image. The
spectral shift is mapped to a false color scale and plotted as hue, while the intensity of the
backscattered light is plotted as saturation, with luminance kept constant. The false color image
thus provides an indication of the spectral shift of the backscattered light as well as its intensity.
A
250 µm
B
250 µm
Figure 21. Spectroscopic OCT imaging. Green and red hue color scale indicates enhanced
optical scattering of short and long wavelengths, respectively. (a) Spectroscopic OCT image of
transition zone between normal squamous epithelium (right) and pathologic columnar epithelium
(left). For normal epithelium, note the relatively smooth color transition from green to red with
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increasing depth in tissue. (b) Spectroscopic OCT image of Barrett’s esophagus. The irregular
epithelial morphology from crypt- and gland-like structures are indicated by areas of red hue
(enhanced long-wavelength scattering) near the surface (arrows).
Figure 21 compares spectroscopic OCT images of normal and Barrett’s esophagus. The
epithelium has a green hue in the false color image, indicating enhance scattering of shorter
wavelengths of light from this layer. Immediately below the epithelium, the lamina propria
contains focal areas of red hue, indicating regions of enhanced longer wavelength scattering.
Since longer wavelengths penetrate more deeply into tissue, it is expected that shorter
wavelengths will be scattered more from structures near the surface while longer wavelengths are
scattered more from deeper structures. Therefore, a shift of hue from green to yellow to red
should occur with depth. Because normal squamous epithelium is relatively homogeneous, this
gradual transition from green to yellow to red is observed in Figure 25 (a). This is in contrast to
Figure 25 (b) where the presence of multiple crypt and glandular structures along the surface
disrupt the morphological structure. The spectroscopic tomogram in Figure 25 (b) shows no
significant degree of uniform green hue (short wavelength scattering) within the epithelium.
Instead, more frequent focal areas of red hue (long wavelength scattering) are present.
Occasional red focal areas are present at the surface as indicated by arrows. Increased
disruption of the normal squamous architectural morphology from the formation of crypts, glands,
and potentially cellular structures such as increased number and size of nuclei, disrupt the normal
spectroscopic scattering behavior in the spectroscopic OCT image.
In conclusion, ultrahigh-resolution OCT techniques achieved 1.1 µm image resolution in in vitro
specimens and showed enhanced resolution of architectural features. Spectroscopic OCT
identified localized regions of wavelength-dependent optical scattering, enhancing the
differentiation of Barrett’s esophagus. These techniques improve structural tissue recognition and
suggest the future potential for resolution and contrast enhancements in clinical studies. We are
currently designing a new catheter compatible with the broadband spectrum of the light source,
which will allow us to perform in vivo endoscopic ultrahigh-resolution OCT and spectroscopic
OCT.
These studies were performed in collaboration with Dr. Jacque van Dam at Stanford Medical
School, Dr. Muthoka Mutinga and Dr. Hiroshi Mashimo at the Veterans Administration Medical
Center.
Ultrahigh Resolution Optical Coherence
Morphology and Tissue Preservation
Sponsors
National Institute of Health
Grant NIH-5-R01-CA75289-04
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-1-R01-AR44812-02
National Institute of Health
Grant NIH-1-R01-HL63953-01
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Tomography
Studies
of
Tissue
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RLE Progress Report 144
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Pei-Lin Hsuing, Paul Herz, Dr. Stephane Bourquin, Dr.
Christian Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Dr. Xingde Li, Dr. Ping Xue, Nirlep
Patel, Karl Schneider, Professor Franz X. Kartner, Professor Mark Brezinski, Dr. Michael
Weinsterin (Brigham and Women’s Hospital), and Professor James G. Fujimoto
To date, many previous studies have compared ex vivo OCT imaging to histopathology. While
some tissues, such as arterial pathology or cartilage, are relatively stable post mortem, others,
such as epithelial tissues, exhibit rapid degradation. It is therefore important to preserve these
tissues with minimal changes in morphology. The goal of this study is to investigate the
difference between in vivo and ex vivo OCT imaging and the effect of different preservation
solutions on image quality using the hamster cheek pouch. The hamster cheek pouch was
chosen because of its easy access and because it is a well established model for carcinogenesis
and cancer progression. The advent of the ultrahigh resolution OCT imaging technology is
important for this study because it enables changes in tissue morphology that may have been
difficult to resolve with standard resolution OCT imaging to be clearly visualized.
OCT imaging was performed using an ultrahigh resolution OCT imaging system [3]. An axial
resolution of 2 µm and a transverse resolution of 5 µm were employed. Several different
preservation solutions were evaluated including: saline at low temperature, saline at room
temperature, phosphate buffered sucrose (PBS140), University of Wisconsin solution (UW), and
10% formalin. Phosphate buffered sucrose and University of Wisconsin solution are common
tissue preservation solutions, while formalin is commonly used for tissue fixation before
histological processing. Following the anesthetization of the animal, the imaging areas in the
cheek pouch were marked with india ink. For each imaging site, a 0.75 mm (axial) by 1 mm
(transverse) OCT image was obtained. Each image consisted of 2000 axial pixels and 1000
transverse pixels to ensure that the pixel density was sufficient for high resolution imaging. The
multiple imaging sites were then resected and the specimens placed in room temperature saline,
chilled saline, PBS140, UW or 10% formalin. The marked areas in the different preservation
solutions were imaged at 30, 50, and 70 minutes post mortem. The image quality, details, and
contrast for each solution were then compared. At the conclusion of the experiments, the tissue
was fixed in formalin and histologically processed.
Figure 22 shows an example of OCT imaging data, which is a compilation of the in vivo images,
the effects of different preservation solutions at 70 minutes postmortem, and the corresponding
histology. The top row (a-e) represents in vivo OCT images of the hamster cheek pouch tissue
taken at different imaging sites. The middle row (f-j) represents ex vivo OCT images of the same
imaging site as the top row taken at 70 minutes postmortem. These areas were imaged ex vivo
while preserved in a saline ice bath (f), room temperature saline (g), phosphate buffered sucrose
140 (h), University of Wisconsin solution (i) and 10% formalin (j). All the OCT images were
collected at a resolution of 2 µm x 5 µm at 800 nm central wavelength. The bottom row
represents the corresponding histology stained with Trichrome. The histology slides are viewed
at 40x and were scaled to match the image size.
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RLE Progress Report 144
Figure 22. In vivo (a-e) and 70 min. postmortem ex vivo (f-j) ultrahigh resolution OCT images of
the hamster cheek pouch tissue under different preservation solutions and their corresponding
histologic cross-sections stained with Trichrome (k-o).
Evaluation of these images indicates that most changes in optical properties and OCT contrast in
tissue occur within minutes after excision for all preservation solutions but formalin. The contrast
between the epithelial cells and the stratum corneum as well as the underlying connective tissue
is rapidly lost in almost all solutions, with the exception of formalin. The low backscattering
nature of the dense connective tissue is not preserved. This could be the result of leakage of
intracellular material from the cells into the extracellular space or expansion of the spaces
between the normally dense collagen fibers. In formalin, the reflection of the two layers above
and below the epithelial cells increases in intensity. Modification of the backscattering intensity of
the lower, looser collagen fibers is also visible ex vivo. The epithelial layer, which is often the
region of interest in neoplastic imaging, appears to be very sensitive to ex vivo preservation
conditions. Although formalin is used routinely for histological processing, it has long been
considered a poor option for tissue preservation in optical imaging studies because it changes the
fluorescent and optical properties of tissue. However, our study indicates that formalin may
actually be a better preservative for ex vivo, morphological OCT imaging. If the goal of the study
is to image and evaluate variations in tissue morphology, then fixation may change the contrast of
some compartments, but it also appears to better maintain the architectural features of interest.
Intravascular Imaging using Optical Coherence Tomography
Sponsors
National Institute of Health
Grant NIH-1-R01-HL63953-01
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-5-R01-CA75289-04
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Pei-Lin Hsiung, Paul Herz, Dr. Stephane Bourquin, Dr.
Christian Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Dr. Ping Xue, Nirlep Patel, Karl
Schneider, Professor Mark Brezinski, and Professor James G. Fujimoto
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RLE Progress Report 144
OCT has been shown to be an effective intravascular imaging technique for the discrimination of
coronary plaques and assessment of interventional procedures such as arterial stent placement
[25, 26]. Guided placement of coronary stents have greatly improved the outcome of
interventional intravascular procedures and constitutes the majority of percutaneous interventions
[27, 28]. Current intracoronary stent procedures require a secondary ultrasound (IVUS) catheter
to be introduced after removal of a balloon catheter in order to assess stent deployment efficacy.
In addition to increased operational time which incurs higher risk to the patient, IVUS resolution is
limited to only ~80 µm at an operational frequency of 40 MHz.
OCT has several advantages for intravascular imaging. High resolution OCT imaging with
discrimination of sub-cellular features has been demonstrated indicating a resolution capability
down to 10-20 µm. Further because OCT is a fiber-optically based imaging technique, extremely
small and compact probes have been demonstrated with outer diameters of <500 µm [13]. The
simple design and components of OCT catheters make them relatively inexpensive as well as
highly compact when compared with IVUS catheters and ultrasound machines. Finally OCT can
be performed at 4-8 frames per second allowing near real-time imaging during coronary
procedures.
In vitro imaging of human coronary arteries was performed using both OCT and IVUS. Figure 23
shows an OCT image of a coronary artery and the corresponding histology. Consistent with
previous work, OCT images of plaque correlate well with histopathology. Most significantly, OCT
images demonstrate relatively good contrast among the layers of the vessel wall. This allows
delineation of both the tunica intima and tunica media within the artery. Histology confirms the
presence and morphology of the intimal hyperplasia seen in the OCT image of this vessel.
Figure 23. A comparison of OCT and histology of a human coronary artery. (a) Discrimination of
tunicae intima, media, and adventitia can be seen in the OCT image. (b) Hematoxylin-eosinstained sections of the artery verify intimal hyperplasia layers seen in the OCT image.
In addition to in vitro arterial imaging, in vivo OCT procedures were done in New Zealand white
rabbits. Images were recorded at 4 frames/second with some slight motion artifacts noted due to
the slower scan speed. Figure 24 (a) shows an aorta image after a manual 2-3 ml/sec saline
flush. In the image, the media and supporting structures are well defined. Backscattering intensity
was high in the media and low in the adventitia which consists primarily of loose connective
tissue. Histology was used to confirm structural morphology seen in the OCT image (Figure 24
(b)).
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RLE Progress Report 144
Figure 24. Aorta image after saline flush and corresponding histology. (a) The media surrounding
the supporting structures are clearly defined. High backscattering is apparent in the media “M”
and low in the adventitia which consists of mainly loose connective tissue. A structure consistent
with the inferior vena cava was also apparent and is imaged through the wall of the aorta. (b)
Corresponding histology to confirm tissue identification. Scale bar represents 500 µm.
The high resolution of OCT allows imaging to be performed near the resolution of histopathology,
offering better discrimination and delineation of arterial structure and morphology. OCT has
demonstrated the potential to have a significant impact on both the identification of high risk
coronary plaques and in guidance of interventional arterial procedures.
Orthopedic Imaging and Detection of Early Osteoarthritic Changes
Sponsors
National Institute of Health
Grant NIH-1-R01-HL63953-01
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-5-R01-CA75289-04
National Institute of Health
Grant NIH-1-R01-AR44812-02
National Institute of Health
Grant NIH-1-R01-EY131780-01
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Pei-Lin Hsuing, Paul Herz, Dr. Stephane Bourquin, Dr.
Christian Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Nirlep Patel, Karl Schneider, Dr. Scott
Martin (Brigham and Women’s Hospital), Professor Mark Brezinski, and Professor James G.
Fujimoto
Osteoarthritis (OA) is the leading cause of chronic disability in developed countries,
symptomatically affecting about 14% of the adult population in the United States alone [29]. The
earliest pathological changes are collagen disorganization, an increase in the water content, and
alteration in glycosaminoglycans (GAGS). Among the later changes are cartilage loss (thinning
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RLE Progress Report 144
effect), fibrillation, and surface erosion [30]. Until recently, osteoarthritis has been accepted as an
inescapable consequence of aging. There is now strong evidence that osteoarthritic progression
can be slowed or modified through surgical and/or pharmacological intervention [31, 32]. Current
imaging technologies however have only a limited ability to monitor changes in articular cartilage.
Available techniques such as plain film and magnetic resonance imaging (MRI) have resolutions
of no better than 10 mm, which limits their use in monitoring cartilage changes [29, 33-35].
Although arthroscopy is widely used in diagnosis of joint disorders [36], it provides magnified
views of only the articular surface. Thus, the pathologic changes under the articular cartilage
surface cannot be visualized. A diagnostic imaging technique capable of high-resolution imaging
of articular cartilage in vivo would be highly valuable to detect disease, to follow its progression
and to monitor therapeutic effectiveness.
In this study, human imaging of normal and osteoarthritic cartilage was demonstrated with optical
coherence tomography using both a hand held probe and an approximately 1 mm arthroscopic
imaging probe. Imaging was performed of human cartilage during at total of 8 total knee
replacements in real time at 4-8 frames per second.
Images compared favorably to
histopathology and clear distinctions were noted between diseased and normal regions. In
addition to identifying cartilage thinning, changes in back-reflection intensity and polarization were
noted in diseased regions.
Figure 25 demonstrates the cross-sectional OCT image of an osteoarthritic joint obtained with the
portable high speed OCT system. The pathologic portion of the cartilage is severely degenerated;
the junction between the normal and pathologic cartilage can be resolved. The figures show a
representative OCT image of the arthritic knee cartilage transitional zone from articular cartilage
to full thickness cartilage loss and to bone (black arrows). The region of the subchondral plate is
demonstrated by red arrows. A corresponding trichrome blue histologic section through the same
image plane as the OCT image is also shown (right).
OCT
Trichrome-Blue
500um
Figure 25. (left) Representative intraoperative OCT image of human osteoarthritic articular
cartilage taken with a hand-held imaging probe during total knee replacement surgery. (right)
Corresponding histological section with trichrome-blue staining.
OCT imaging was also performed in human knees with an OCT arthroscope. Again, highresolution structural assessments of articular cartilage could be obtained, including the
demonstration of polarization sensitivity. Imaging was performed using a 1 mm diameter optical
imaging catheter. The catheter is shown schematically in Figure 6 and consists of a single mode
fiber in a rotating speedometer cable with a distal lens and microprism encased in a transparent
housing [3]. Imaging was performed in vitro on an intact human knee obtained from autopsy to
examine the feasibility of performing OCT through a cannula during arthroscopy.
Figure 26 shows examples of images obtained with the OCT imaging probe. The images show
the boundary of cartilage and bone (left) and a ligament (right). The ligament has a high
birefringence due to the organization of collagen. The multiple striped features in the OCT image
are the result of polarization birefringence.
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RLE Progress Report 144
500 um
500 um
Figure 26. Representative OCT images taken using an OCT imaging catheter introduced through
a cannula in an intact knee in vitro. The image on the left shows the junction between cartilage
and bone. The image on the right shows a ligament which exhibits strong birefringence.
These studies demonstrate that OCT imaging can be performed in an intact knee joint in vitro
using an OCT imaging catheter introduced through a cannula in a minimally invasive procedure.
This procedure is an important step toward performing OCT imaging in the clinic. It is clear that
OCT represents a promising new technology for monitoring changes in articular cartilage, both in
an open surgical field and through an arthroscope.
Prostate Cancer Detection and Imaging Neoplastic Changes
Sponsors
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
Air Force Office of Scientific Research
Grant F49620-98-01-0084
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-5-R01-CA75289-04
National Institute of Health
Grant NIH-1-R01-EY131780-01
Binney Radiation Oncology Fund
Project Staff
Aaron Aguirre, Tony Ko, Pei-Lin Hsuing, Paul Herz, Dr. Stephane Bourquin, Dr. Christian
Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Dr. Christopher Kroeger, Dr. Xingde Li, Karl
Schneider, Dr. Anthony D’Amico, Dr. Michael Weinstein, Dr. Jerome Richie (Brigham and
Women’s Hospital), Dr. Phillip Kantoff (Dana Farber), and Professor James G. Fujimoto
The lifetime risk in Western society of developing histologic evidence of prostate cancer has been
estimated to be 30%, with the risk of developing clinical disease 10% and the probability of
mortality due to the disease roughly 3% [37]. Core biopsy is the standard method for diagnosing
prostate cancer. However biopsy techniques often have sampling errors, which lead to false
negatives. There has therefore been significant interest in novel techniques to improve the
diagnostic accuracy in prostate cancer. Ex vivo, ultrahigh resolution OCT imaging studies
suggest that differentiation of glandular architectural morphology associated with prostate cancer
is feasible [38]. Devices such as OCT needles have been demonstrated which enable imaging in
solid tissues or organs [13]. A technology capable of imaging the prostate in vivo at high
resolution could be integrated with core biopsy to reduce sampling error.
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RLE Progress Report 144
We have performed preliminary imaging studies to investigate the feasibility of OCT for imaging
prostate pathology in vivo. Imaging intraoperatively was chosen to eliminate the potential effects
of tissue degradation on contrast or image quality as well as to investigate the feasibility of using
OCT in an intraoperative setting. Informed consent was obtained from patients with prostate
cancer scheduled to undergo nerve-sparing radical prostatectomy. The patient selection criterion
was clinical stage T1-T4 adenocarcinoma of the prostate as evident on core biopsy. This patient
population was selected because of the likelihood of a large tumor that might be visible with
intraoperative OCT imaging through the prostatic capsule.
To date, imaging has been performed on 2 patients with clinical stage T1-T4 adenocarcinoma of
the prostate undergoing nerve-sparing radical prostatectomy. Images were taken of the posterior
surface of the prostate after the urethra was divided and the prostate is retracted. Figure 27
shows an example intraoperative OCT image of the right posterior peripheral zone of the prostate
with corresponding histology. Striations in the fibromuscular stroma are clearly visible in both the
OCT image and histology.
250 µm
250 µm
Figure 27. Representative intraoperative OCT image of normal prostate in vivo. Striations in the
fibromuscular stroma are visible in both the OCT image and in the corresponding histologic
photomicrograph.
A
250 µm
B
250 µm C
250 µm
Figure 28. Representative in vivo OCT images of (a) large vein just underneath the prostate
capsule, (b) a large gland, (c) lipid layer overlying normal prostate tissue.
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RLE Progress Report 144
A variety of structures in the normal prostate can also be identified with OCT. Figure 28 shows
examples of a large vein underlying the prostate capsule, a large benign gland, and a thin layer of
lipid-filled structures overlying normal prostatic tissue. Imaging was also performed ex vivo in the
pathology laboratory. Imaging in the pathology laboratory immediately after surgical excision
enables accurate registration of OCT images with histology as well as a comparison of in vivo
versus ex vivo imaging. In order to preserve diagnostic integrity, specimens often cannot be
removed from the hospital, and thus the capability of imaging in the pathology lab setting is
especially important as a research tool. OCT images of the prostate taken ex vivo showed no
significant degradation in contrast or image quality when imaging is performed immediately after
excision of the specimen.
These studies demonstrate the feasibility of using OCT to image human prostatic tissue in vivo.
Additional studies are required to evaluate the ability of OCT to differentiate prostate cancer.
Advances in solid-state laser and nonlinear fiber technology have allowed development of
ultrahigh resolution OCT which enables 1-3 µm resolution imaging, and spectroscopicallyresolved OCT using broadband light sources also promises to improve tissue differentiation and
image contrast. The ability to visualize tissue pathology in situ and in real time promises to
improve both diagnosis and therapy of prostate carcinoma.
These studies were conducted in collaboration with Dr. Mike Weinstein and Dr. Anthony D’Amico
at Brigham and Women’s Hospital and Dr. Phillip Kantoff at the Dana Farber Cancer Center.
Surgical Guidance using Optical Coherence Tomography
Sponsors
Air Force Office of Scientific Research (MFEL)
Grant F49620-01-1-0186
Air Force Office of Scientific Research
Grant F49620-98-01-0084
National Institute of Health
Grant NIH-2-R01 EY11289-15
National Institute of Health
Grant NIH-5-R01-CA75289-04
National Institute of Health
Grant NIH-1-R01-AR44812-02
National Institute of Health
Grant NIH-1-R01-HL63953-01
Project Staff
Aaron Aguirre, Ravi Ghanta, Tony Ko, Pei-Lin Hsiung, Paul Herz, Dr. Stephane Bourquin, Dr.
Christian Chudoba, Dr. Wolfgang Drexler, Dr. Ingmar Hartl, Dr. Christopher Kroeger, Dr. Ping
Xue, Nirlep Patel, Karl Schneider, Dr. Anthony D’Amico, Dr. Michael Weinstein, Dr. Jerome
Richie (Brigham and Women’s Hospital), Dr. Phillip Kantoff (Dana Farber), Professor Mark
Brezinski, and Professor James G. Fujimoto
Optical coherence tomography can image microstructure in tissue in situ and in real time with
image resolutions in the 1-10 µm range [1, 3, 39]. Previous studies have demonstrated that
changes in tissue architectural morphology associated with neoplasia can be identified [39-41],
and recent advances such as ultrahigh resolution and spectroscopic OCT promise to further
improve the differentiation of tissue pathology [3, 9]. The development of intraoperative OCT
enables pathology to be visualized in situ and in real time in tissues that were previously
inaccessible in vivo. Combining Intraoperative studies with imaging in the pathology laboratory
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RLE Progress Report 144
immediately after surgical excision enables accurate registration of OCT images with histology as
well as a comparison of in vivo versus ex vivo imaging. In order to preserve diagnostic integrity,
specimens often cannot be removed from the hospital, and thus the capability of imaging in the
pathology lab setting is especially important as a research tool. Techniques that can visualize
microscopic disease in situ and in real time promise to improve diagnostic accuracy. The ability
to visualize tissue structures intraoperatively should also facilitate the guidance of microsurgical
procedures [42].
In order to perform intraoperative imaging, an OCT system must be ergonomically well-designed
and robust to function under the tight confines and time constraints of the surgical suite. An
image delivery device that is sterilizable using standard hospital sterilization protocols is also
required. We have developed a sterilizable hand-held surgical probe used with a portable OCT
system to image in the surgical suite and pathology laboratory. The imaging probe has a 2 cm
working distance which allows the surgeon to maneuver the beam within the tight confines of the
surgical field. Figure 29 shows a schematic of the probe and a photograph of the operative field
during radical prostatectomy. A low power green aiming beam allows the surgeon to localize the
imaging area and facilitate marking for correlation with histology. The probe has a stainless steel
outer tube with a transparent window which can be sterilized using standard hospital sterilization
procedures and slipped over the distal end of the probe prior to imaging. The proximal end of the
probe including cables was enclosed in a sterile surgical camera bag. The probe produced a
focused spot with an 18 µm transverse resolution. The portable OCT imaging system used a
semiconductor amplifier (AFC) low coherence light source operating at 1300 nm wavelength
having an axial resolution of 15 µm in air. Images of 500 to 250 transverse pixels by 700 axial
pixels could be obtained at 4-8 frames per second, respectively, with 95 dB sensitivity using a
less than 6 mW incident power imaging beam.
Urethra
l
f2
f2 + f3
f3
f4
Prosta
f1
Semin
al
fiber
Figure 29. (left) Schematic OCT hand-held probe and (right) picture of the surgical field. The
probe uses a galvanometer controlled mirror and Hopkin’s lens relay system to acquire OCT
images over a 3 mm field of view. A sterile stainless steel outer sheath covers the distal end of
the probe. A urethral catheter aids intraoperative manipulation of the prostate.
We have performed intraoperative imaging studies of patients with prostate cancer undergoing
radical prostatectomy and patients with osteoarthritis undergoing open field surgery. Studies were
performed to investigate the feasibility of using OCT for imaging these pathologies in vivo as well
as to investigate OCT intraoperative imaging in a more general context. Imaging in the pathology
laboratory has also been performed immediately after surgical excision to enable accurate
registration of OCT images with histology as well as a comparison of in vivo versus ex vivo
imaging. Preliminary results demonstrate the feasibility of intraoperative OCT imaging in human
subjects as well as imaging in a pathology laboratory environment. Recently we have designed a
version of the probe that uses two orthogonally positioned galvanometer mirrors to enable
scanning of the OCT beam in the en face plane. This has enabled us to use the probe for in vivo
investigations of optical coherence microscopy, a technique that combines standard OCT imaging
with confocal microscopy. Extension of these technologies into the intraoperative setting and the
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RLE Progress Report 144
pathology laboratory should enable comprehensive imaging studies. The ability to visualize tissue
pathology in situ and in real time promises to improve both diagnosis and therapy.
Publications
U. Morgner, W. Drexler, F.X. Kärtner, X. D. Li, C. Pitris, E. P. Ippen, J. G. Fujimoto,
“Spectroscopic optical coherence tomography,” Optics Letters 25, 111-113, January 15, 2000.
J. G. Fujimoto, C. Pitris, S. Boppart, and M. Brezinski, “Optical coherence tomography, an
emerging technology for biomedical imaging and optical biopsy,” Neoplasia 2, 9-25, January
2000.
C. Pitris, C. Jesser, S. A. Boppart, D. Stamper, M. E. Brezinski, and J. G. Fujimoto, “Feasibility of
optical coherence tomography for high resolution imaging of human gastrointestinal tract
malignancies,” J. Gastroenterology 35, 87-92, February 2000.
P. Patwari, N. J. Weissman, S. A. Boppart, C. A. Jesser, D. Stamper, J. G. Fujimoto, and M. E.
Brezinski, “Assessment of coronary plaque with optical coherence tomography and high
frequency ultrasound,” Journal of the American College of Cardiology, 85, 641-644, March 2000.
S. A. Boppart, J. M. Herrmann, and J. G. Fujimoto, “Optical coherence tomography: Applications
in surgical diagnosis, guidance and intervention,” Medical Imaging International 10, 14-20, MarchApril 2000.
A. V. D’Amico, M. Weinstein, X. Li, J. P. Richie, and J. G. Fujimoto, “Optical coherence
tomography as a method for identifying benign and malignant microscopic structures in the
prostate gland,” Urology 55, 783-787, 2000.
J. Van Dam and J. G. Fujimoto, “Imaging beyond the endoscope,” Gastrointestinal Endoscopy
51, Part 1, 512-516, 2000.
W. Drexler, U. Morgner, R. K. Ghanta, J. S. Schuman, F. X. Kartner, M. R. Hee, E. P. Ippen, and
J. G. Fujimoto, “New technology for ultrahigh resolution optical coherence tomography of the
retina,” in The Shape of Glaucoma, Quantitative Neural Imaging Techniques, H. G. Lemij and J.
S. Schuman, Eds., pp. 75-104, Kugler Publications, The Hague, Netherlands, 2000.
X. D. Li, C. Chudoba, T. Ko, C. Pitris, and J.G. Fujimoto, “Imaging needle for optical coherence
tomography,” Optics Letters 25, October 2000.
X. D. Li, S. A. Boppart, J. Van Dam, H. Mashimo, M. Mutinga, W. Drexler, M. Klein, C. Pitris, M.
L. Krinsky, M. E. Brezinski, and J. G. Fujimoto, “Optical coherence tomography: advanced
technology for the endoscopic imaging of Barrett's Esophagus,” Endoscopy 32, 921-30,
December 2000.
W. Drexler, U. Morgner, R. K. Ghanta, F. X. Kärtner, J. S. Schuman, J.G. Fujimoto, “Ultrahigh
resolution ophthalmic optical coherence tomography,” Nature Medicine 7, 502-507, April 2001.
C. Chudoba, J. G. Fujimoto, E. P. Ippen, H. A. Haus, U. Morgner, F. X. Kaertner, V. Scheuer, G.
Angelow, and T. Tschudi, “All-solid-state Cr:forsterite laser generating 14 fs pulses at 1.3 mm,”
Optics Letters 26, 292-294, March 2001.
I. Hartl, X.D. Li, C. Chudoba, R. Ghanta, T. Ko, J.G. Fujimoto, J. K. Ranka, R. S. Windeler, and A.
J. Stentz, “Ultrahigh resolution optical coherence tomography using continuum generation in an
air-silica microstructure optical fiber,” Optics Letters 26, 608-610, May 2001.
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RLE Progress Report 144
C. Pitris, K. T. Saunders, J. G. Fujimoto, and M. E. Brezinski, “High resolution imaging of the
middle ear with optical coherence tomography, a feasibility study,” Archives of Otolaryngology
Head and Neck Surgery, 127, 637-42, June 2001.
W. Drexler, D. Stamper, C. Jesser, X. D. Li, C. Pitris, K. Saunders, S. Martin, M. B. Lodge, J. G.
Fujimoto, and M. E. Brezinski, “Correlation of collagen organization with polarization sensitive
imaging of in vitro cartilage: Implications for osteoarthritis,” Journal of Rheumatology, 28, 1311-8,
June 2001.
S. A. Boppart, J. M. Herrmann, C. Pitris, D. L. Stamper, M. E. Brezinski, and J. G. Fujimoto,
“Real-time optical coherence tomography for minimally-invasive imaging of prostate ablation,”
Computer Aided Surgery 6, 94-103, 2001.
X. Li, T.H. Ko, J.G. Fujimoto, “Intraluminal fiber-optic Doppler imaging catheter for structural and
functional optical coherence tomography,” Optics Letters, 26, 1906-09, December 2001.
J.G. Fujimoto, “Optical coherence tomography,” (C. R. Acad. Sci. Paris, Serie IV) Applied
Physics, 1099-1111, 2001.
X. Li, S. Martin, C. Pitris, R. Ghanta, C. Jesser, M. Spectro, J. G. Fujimoto, and M. E. Brezinski,
“High-resolution imaging of osteoarthritic cartilage,” Journal of Rheumatology, submitted for
publication.
X. D. Li, H. Gold, N. Weissman, K. Saunders, C. Pitris, J. G. Fujimoto, and M. E. Brezinski,
“Optical Coherence tomography guided stent deployment: a comparison with IVUS in an in vivo
rabbit model,” Journal of the American College Cardiology, submitted for publication.
Abstracts and Conference Presentations
I. Hartl, X. Li, C. Chudoba, R. K. Ghanta, T. H. Ko, J. G. Fujimoto, J. K. Ranka, R. S. Windeler,
and A. J. Stentz, “Ultrahigh resolution OCT using continuum generation in an air-silica
microstructure optical fiber,” International Biomedical Optics Symposium, BIOS’2001, San Jose,
CA, January 20-26, 2001, paper 4251-07.
X. Li, C. Chudoba, T. H. Ko, C. Pitris, R. K. Ghanta, and J. G. Fujimoto, “Imaging solid tissues
with an OCT imaging needle,” International Biomedical Optics Symposium, BIOS’2001, San Jose,
CA, January 20-26, 2001, paper 4251-09.
P. Hsiung, X. Li, I. Hartl, C. Chudoba, T. H. Ko, R. K. Ghanta, and J. G. Fujimoto, “High-speed
path length scanning for optical coherence tomography,” International Biomedical Optics
Symposium, BIOS’2001, San Jose, CA, January 20-26, 2001, paper 4251-15.
C. Pitris, T. H. Ko, W. Drexler, R. K. Ghanta, X. Li, C. Chudoba, I. Hartl, J. G. Fujimoto, and M. E.
Weinstein, “Ultrahigh resolution in-vivo OCT imaging and tissue preservation,” International
Biomedical Optics Symposium, BIOS’2001, San Jose, CA, January 20-26, 2001, paper 4251-29.
W. Drexler, R. K. Ghanta, J. S. Schuman, T. H. Ko, U. Morgner, F. X. Kärtner, and J. G. Fujimoto,
“In-vivo optical biopsy of the human retina using optical coherence tomography,” International
Biomedical Optics Symposium, BIOS’2001, San Jose, CA, January 20-26, 2001, paper 4251-32.
P. Hsiung, C. Pitris, X. Li, A. Goodman, R. K. Ghanta, T. H. Ko, C. Chudoba, I. Hartl, W. Drexler,
M. E. Brezinski, and J. G. Fujimoto, “In-vivo cervical imaging with an integrated optical coherence
tomography colposcope,” International Biomedical Optics Symposium, BIOS’2001, San Jose, CA,
January 20-26, 2001, paper 4254-31.
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X. Li, S. A. Boppart, J. Van Dam, H. Mashimo, M. Mitinga, W. Drexler, M. Klein, C. Pitris, M. L.
Krinsky, M. E. Brezinski, and J. G. Fujimoto, “Imaging Barrett’s esophagus with optical coherence
tomography,” International Biomedical Optics Symposium, BIOS’2001, San Jose, CA, January
20-26, 2001, paper 4254-32.
T.H. Ko, R.K. Ghanta, I. Hartl, W. Drexler, J.G. Fujimoto, “Ultrahigh resolution optical coherence
tomography for quantitative topographic mapping of retinal and intraretinal architectural
morphology,” Association for Research in Vision and Ophthalmology Annual Meeting, Fort
Lauderdale, FL, April 29-May 4, 2001, paper 4251-B270.
I. Hartl, T. Ko, R.K. Ghanta, W. Drexler, A. Clermont, S.E. Bursell, J.G. Fujimoto, “In vivo
ultrahigh resolution optical coherence tomography for the quantification of retinal structure in
normal and transgenic mice,” Association for Research in Vision and Ophthalmology Annual
Meeting, Fort Lauderdale, FL, April 29-May 4, 2001, paper 4252-B271.
W. Drexler, I. Hartl, T. Ko, R.K. Ghanta, J.S. Shuman, S.E. Bursell, J.G. Fujimoto, “Ultrahigh
resolution imaging of retinal diseases with optical coherence tomography,” Association for
Research in Vision and Ophthalmology Annual Meeting, Fort Lauderdale, FL, April 29-May 4,
2001, paper 636.
P. Hsiung, C. Chudoba, C. Pitris, X. Li, T. Ko, and J. G. Fujimoto, “In vivo colposcopie imaging of
neoplastic tissue using optical coherence tomography,” Conference on Laser and Electro-Optics,
CLEO’01, Baltimore, MD, May 6-11, 2001, paper CTuY2.
T. Ko, C. Pitris, I. Hartl, R. Ghanta, C. Chudoba, X. D. Li, W. Drexler, J. G. Fujimoto, and M.
Weinstein, “Ultrahigh resolution in vivo versus ex vivo OCT imaging and tissue preservation,”
Conference on Laser and Electro-Optics, CLEO’01, Baltimore, MD, May 6-11, 2001, paper
CTuY2.
I. Hartl, C. Chudoba, T. Ko, X. D. Li, J. G. Fujimoto, M. Brezinski, and R. S. Windeler, “Imaging
water absorption with spectroscopic optical coherence tomography,” Conference on Laser and
Electro-Optics, CLEO’01, Baltimore, MD, May 6-11, 2001, paper CWE2.
P. Hsiung, X. D. Li, I. Hartl, T. Ko, C. Chudoba, and J. G. Fujimoto, “High-speed path length
scanning using a Herriott cell delay line,” Conference on Laser and Electro-Optics, CLEO’01,
Baltimore, MD, May 6-11, 2001, poster CWA62.
J. G. Fujimoto, “Biomedical imaging and optical biopsy using optical coherence tomography,”
24th Annual Meeting of the Japan Society of Bronchology, May 24-25, 2001, Chiba, Japan,
invited talk.
J. G. Fujimoto, “Optical coherence tomography: current status and future prospects,” Schepens
International Society Meeting, 2001 and Beyond- A Vetreoretinal Odyssey, Las Vegas, NV, June
6-9, 2001, invited talk.
I. Hartl, C. Chudoba, T. Ko, X. D. Li and J. G. Fujimoto, R. S. Windeler, “Optical coherence
tomography using continuum generation in an air-silica microstructure optical fiber,” European
Conference on Biomedical Optics, ECBO, Munich, June 17-22, 2001, paper MH2.
J. G. Fujimoto, I. Hartl, W. Drexler, C. Chudoba, T. Ko, X. Li, P. Hsiung, U. Morgner, and F. X.
Kärtner, “Biomedical imaging using ultralight resolution optical coherence tomography,” XVII
International Conference on Coherent and Nonlinear Optics (ICONO-2001), Minsk, Belarus, June
26 - July 1, 2001, paper 716, invited talk.
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J. G. Fujimoto, “Optical biopsy using optical coherence tomography,” Engineering Foundation
Conference, Advances In Optics For Biotechnology, Medicine And Surgery, Banff, Canada, July
22-27, 2001.
J. G. Fujimoto, “Optical coherence tomography imaging:
technology and applications,”
International Congress on Analytical Sciences 2001, Tokyo, Japan, August 6-10, 2001, invited
talk.
X. Li, I. Hartl, C. Chudoba, W. Drexler, T. Ko, P. Hsiung, C. Pitris, R. Ghanta, F. X. Kärtner, U.
Morgner, M. E. Brezinski, and J. G. Fujimoto, “Ultrahigh resolution and spectroscopic optical
coherence tomography,” OSA Annual Meeting, ILS-XVII: 17th Interdisciplinary Laser Science
Conference, Long Beach, CA, October 14-18, 2001, paper WV1, invited talk.
J. G. Fujimoto, “Optical coherence tomography for structural and functional imaging,” Imaging in
2020, Jackson Hole, WY, September 30 – October 4, 2001, invited talk.
W. Drexler, I. Hartl, T. H. Ko, R. K. Ghanta, H. Sattmann, B. Hermann, B. Povazay, K. K.
Bizheva, J. S. Schuman, O. Findl, S. Bursell, M. Stur, A. F. Fercher, and J. G. Fujimoto,
“Enhanced visualization of retinal pathologies using ultrahigh resolution optical coherence
tomography,” Photonics West, SPIE, San Jose, CA, January 19-25, 2002, paper 4611-41.
X. Li, T. H. Ko, and J. G. Fujimoto, “Intraluminal fiber optic Doppler imaging catheter for structural
and functional optical coherence tomography,” Photonics West, SPIE, San Jose, CA, January 1925, 2002, paper 4619-22.
T. H. Ko, I. Hartl, W. Drexler, R. K. Ghanta, and J. G. Fujimoto, “Ultrahigh resolution optical
coherence tomography for quantitative topographic mapping of retinal and intraretinal
architectural morphology,” Photonics West, SPIE, San Jose, CA, January 19-25, 2002, paper
4619-40.
J. G. Fujimoto, I. Hartl, P. Xue, T. H. Ko, C. Chudoba, W. J. Wadsworth, T. A. Birks, and R. S.
Windeler, “Supercontinuum generation in photonic crystal fibers as light sources for OCT,”
Photonics West, SPIE, San Jose, CA, January 19-25, 2002, paper 4633-26, invited talk.
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U. Morgner, W. Drexler, X. Li, F. X. Kaertner, C. Pitris, S. A. Boppart, E. P. Ippen, and J.
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J. M. Schmitt, S. H. Xiang, and K. M. Yung, “Differential absorption imaging with optical
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U. S. Sathyam, J. B. W. Colston, L. B. D. Silva, and M. J. Everett, “Evaluation of optical
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X. Li, T. H. Ko, and J. G. Fujimoto, “Intraluminal fiber-optic Doppler imaging catheter for
structural and functional optical coherence tomography,” Optics Letters, vol. 26, pp.
1906-1908, 2001.
X. D. Li, C. Chudoba, T. Ko, C. Pitris, and J. G. Fujimoto, “Imaging needle for optical
coherence tomography,” Optics Letters, vol. 25, 2000.
M. Rajadhyaksha, M. Grossman, D. Esterowitz, R. H. Webb, and R. R. Anderson, “In vivo
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J. A. Izatt, M. R. Hee, G. M. Owen, E. A. Swanson, and J. G. Fujimoto, “Optical
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C. A. Puliafito, M. R. Hee, C. P. Lin, E. Reichel, J. S. Schuman, J. S. Duker, J. A. Izatt, E.
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J. S. Schuman, T. Pedut-Kloizman, E. Hertzmark, M. R. Hee, J. R. Wilkins, J. G. Coker,
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X. D. Li, S. A. Boppart, J. Van Dam, H. Mashimo, M. Mutinga, W. Drexler, M. Klein, C.
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P. Patwari, N. J. Weissman, S. A. Boppart, C. A. Jesser, D. Stamper, J. G. Fujimoto, and
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A. V. D'Amico, M. Weinstein, X. Li, J. P. Richie, and J. Fujimoto, “Optical coherence
tomography as a method for identifying benign and malignant microscopic structures in
the prostate gland,” Urology, vol. 55, pp. 783-7, 2000.
J. G. Fujimoto, M. E. Brezinski, G. J. Tearney, S. A. Boppart, B. Bouma, M. R. Hee, J. F.
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tomography,” Nature Medicine, vol. 1, pp. 970-972, 1995.
C. Pitris, A. Goodman, S. A. Boppart, J. J. Libus, J. G. Fujimoto, and M. E. Brezinski,
“High-resolution imaging of gynecologic neoplasms using optical coherence tomography,”
Obstetrics and Gynecology, vol. 93, pp. 135-9, 1999.
A. M. Sergeev, V. M. Gelikonov, G. V. Gelikonov, F. I. Feldchtein, R. V. Kuranov, N. D.
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OCT imaging of precancer and cancer states of human mucosa,” Optics Express, vol. 1,
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S. A. Boppart, B. E. Bouma, C. Pitris, G. J. Tearney, J. F. Southern, M. E. Brezinski, and
J. G. Fujimoto, “Intraoperative assessment of microsurgery with three-dimensional optical
coherence tomography,” Radiology, vol. 208, pp. 81-86, 1998.
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