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Title
High-speed fiber based polarization-sensitive optical coherence tomography of in vivo
human skin.
Permalink
https://escholarship.org/uc/item/01h897jj
Journal
Optics letters, 25(18)
ISSN
0146-9592
Authors
Saxer, CE
de Boer, JF
Park, BH
et al.
Publication Date
2000-09-01
DOI
10.1364/ol.25.001355
Copyright Information
This work is made available under the terms of a Creative Commons Attribution License,
availalbe at https://creativecommons.org/licenses/by/4.0/
Peer reviewed
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September 15, 2000 / Vol. 25, No. 18 / OPTICS LETTERS 1355
High-speed fiber–based polarization-sensitive optical
coherence tomography of in vivo human skin
Christopher E. Saxer, Johannes F. de Boer, B. Hyle Park, Yonghua Zhao, Zhongping Chen, and J. Stuart Nelson
Beckman Laser Institute and Medical Clinic, University of California, Irvine, Irvine, California 92612
Received March 16, 2000
A high-speed single-mode fiber – based polarization-sensitive optical coherence tomography (PS OCT) system
was developed. With a polarization modulator, Stokes parameters of reflected f light for four input polariza-
tion states are measured as a function of depth. A phase modulator in the reference arm of a Michelson
interferometer permits independent control of the axial scan rate and carrier frequency.
In vivo PS OCT
images of human skin are presented, showing subsurface structures that are not discernible in conventional
OCT images. A phase retardation image in tissue is calculated based on the ref lected Stokes parameters of
the four input polarization states. © 2000 Optical Society of America
OCIS codes: 170.0170, 170.4500, 260.1440, 260.5430, 110.7050, 170.1870.
Optical coherence tomography (OCT) is an emerging
technology for noninvasive imaging of biological tissue.
Based on a Michelson interferometer, the technique
measures spatially resolved ref lected intensity
in tissue, offering a dynamic range in excess of
100 dB.
1
The development of polarization-sensitive
optical coherence tomography (PS OCT) has per-
mitted additional information on the polarization
properties of tissue carried by the ref lected light
to be obtained.
2–6
Many biological tissues such as
tendon, muscle, nerve, bone, cartilage, and teeth
exhibit birefringence because of their linear or fibrous
structure.
7
The advantages of PS OCT are enhanced
contrast and specif icity in identifying structures in
OCT images by detection of changes induced in the
polarization state of light that ref lected from the
sample.
Previous PS OCT systems were air-spaced inter-
ferometers that used bulk optical components that
permitted precise control over the polarization state
of light in the sample and reference arms.
2–6
Fiber-
based interferometers offer distinct advantages in
terms of system alignment and handling but pose
design problems owing to polarization changes in-
duced in optical fibers. Polarization-maintaining
(PM) f ibers have large birefringence with a beat
length of 2–3 mm. The energy of wave components
along the primary axes of the fiber is preserved, but
the relative phase is lost owing to the difference in
experienced optical path length. To determine the
Stokes parameters, such phase information is needed.
5
Single-mode (SM) f ibers have polarization-mode
dispersion (PMD). Because of the random birefrin-
gence induced by core ellipticity and noncircularly
symmetric stresses, SM fibers propagate two nearly
degenerate orthogonal polarization states. Differen-
tial phase delay (DPD) and differential group delay
between these two states cause, respectively, an
evolution of the polarization states along the fiber and
a broadening of the interferogram in an OCT system.
For fiber length
L shorter than mode-coupling length
h, DPD and differential group delay are directly pro-
portional to fiber length. This relationship changes
into a square-root dependence for L .. h, indicating
the underlying one-dimensional random-walk nature
of PMD.
8
We used a single-mode fiber (Corning
SMF-28) with a PMD # 0.1 ps兾
p
km if L . h and
#0.1
p
km兾
p
h ps兾km if L , h, which yields an optical
path-length difference between orthogonal polariza-
tion states that is in either case less than 2 mmfor
4.4 m of f iber.
In Fig. 1 a single-mode fiber–based PS OCT system
is presented. We minimized the total PMD by using
a short f iber length in the sample arm (1.75 m) and
in the reference arm (45 cm SM; 75 cm PM). A
low-coherence source (AFC Technologies) with a
FWHM bandwidth of 80 nm centered at 1310 nm was
polarized by a bulk polarizer and coupled back into
the fiber. Quarter- and half-wave plates before the
polarizer were used to select the polarization state of
the source with the highest power (8 mW). Quarter-
and half-wave plates after the polarizer prepared
the polarization such that after a short fiber length
(15 cm) the light emerged with equal-magnitude wave
components parallel and perpendicular to the optic
axis of a bulk electro-optic polarization modulator
(New Focus 4104). The modulator permits control of
the polarization state over a grand circle on Poincaré’s
sphere, as shown in Fig. 2. The dashed grand circle
indicates the possible polarization states immediately
Fig. 1. Schematic of the fiber-based PS OCT system.
Source: 1310 nm, 80-nm bandwidth; FB, fiber bench
with bulk polarizer and quarter- and half-wave plates;
Pol. Mod., LiNbO
3
polarization modulator mounted in
fiber bench; PC’s, static polarization controllers; PB,
polarizing beam splitter; PD, photo diodes; Phase Mod.,
phase modulator; RSOD, defined in text; HP, hand piece
mounted upon a motorized linear translation stage.
Driving waveforms to the galvo of the RSOD and the
polarization modulator show their respective timings.
0146-9592/00/181355-03$15.00/0 © 2000 Optical Society of America
1356 OPTICS LETTERS / Vol. 25, No. 18 / September 15, 2000
Fig. 2. Poincaré’s sphere, with the Q, U , and V axes indi-
cated. The dashed grand circle and axes show the possible
polarization states after the polarization modulator. The
dotted grand circle and axes show a realization of the po-
larization states at the sample arm fiber tip as a result of
PMD. The angles between the Stokes vectors are main-
tained in the absence of polarization-dependent loss.
after the modulator. A four-step driving function,
in which each step introduces a p兾2 phase shift,
cycles the light over four Stokes vectors, indicated
by the dashed axes in the grand circle, before a
fiber 2 3 2 coupler. In the reference arm, a static
polarization controller is aligned such that for all four
polarization states half of the light is transmitted
through a PM fiber pigtailed phase modulator (JDS
Uniphase), which by its structure polarizes the light.
The PM fiber is also used to couple the light into a
rapid-scanning optical delay line (RSOD),
9,10
which
is operated with the spectrum centered on the galvo
mirror. The RSOD thus generates only a group delay
and no phase delay; the carrier of the interferometric
signal at the detector is generated only by the phase
modulator. The phase modulator is driven by a saw-
tooth waveform at 1 MHz, generating a maximum 2p
phase shift after double passage. The sample arm
consists of a fiber with a collimator and a focusing lens
mounted upon a motorized linear translation stage.
As a result of DPD, the polarization state at the tip
of the sample arm fiber is unknown. However, in
a lossless fiber (with a total differential group delay
smaller than the coherence time of the light), the
transformations in Poincaré’s sphere are orthonormal,
preserving the angles between the four Stokes vectors.
The dotted grand circle and axes in Fig. 2 indicate a
possible realization of the four polarization states at
the fiber tip. In the detection arm a static polariza-
tion controller before the polarizing beam splitter is
aligned such that the light from the reference arm is
split equally over both detectors. Electronic signals
are high-pass filtered, amplified, and digitized by
a 12-bit dual-channel 10-Msamples兾s per channel
analog–digital board (Gage Applied Sciences Inc.).
Data processing consists of lock-in detection in soft-
ware of the sine and cosine components at the reference
frequency of 1 MHz. The sine and cosine components
of 5-mm segments in each A line (depth profile) of 2-mm
length are processed to yield the Stokes parameters as
described earlier
5,7
:
I 苷 sin
H
2
1 cos
H
2
1 sin
V
2
1 cos
V
2
, (1)
Q 苷 sin
H
2
1 cos
H
2
2 sin
V
2
2 cos
V
2
, (2)
U 苷 2 sin
H
3 sin
V
1 2 cos
H
3 cos
V
, (3)
V 苷 2 sin
H
3 cos
V
2 2 cos
H
3 sin
V
, (4)
where sin
H, V
and cos
H, V
are the sine and cosine com-
ponents; subscripts H and V indicate, respectively, the
horizontal and vertical polarization channels; and I, Q,
U , and V are the Stokes parameters.
Figure 3 shows a set of four images (I, Q, U , and V )
of human skin from the finger of a volunteer subject
for a single-input polarization state. The finger was
pressed against a glass slide with index-matching f luid
to reduce the surface ref lection from the tissue. Ob-
vious changes are visible in Stokes images Q, U , and
V , which indicate the presence of structures that are
not seen in ref lectivity image I that are changing the
polarization state of the light.
Figure 4 shows the results of using a polarization
modulator to image human skin from the forearm of a
volunteer subject. Modulating the incident light over
four polarization states generates 16 images, 4 for each
input state. The RSOD was driven by a 624-Hz trian-
gular waveform, and the polarization modulator was
driven by a 312-Hz four-step waveform. The veloc-
ity of the linear translation stage was 1.56 mm兾s, and
the scan took 2 s. Figure 4 shows a set of two im-
ages (A and B) containing the information from these
16 images.
We constructed the first image (A) by averaging the
ref lectivity images of the four different input states.
The second image was constructed as follows: The
four different input states lay along perpendicular axes
on a grand circle in Poincaré’s sphere; two pairs of
input states were collinear but in opposite directions.
Their Stokes parameters Q, U , and V differed by only
a minus. The Stokes parameter images Q, U , and V
for such an input pair were very similar, except for the
minus. For each pair we averaged the I images by
addition and the Q, U , and V images by subtraction.
The resulting eight images (one set of I , Q, U, and V
images per input pair) define two Stokes vectors that
are described by lengths I
1
and I
2
and three component
(Q, U , and V ) vectors
S
1
and S
2
with unit length.
The second image (B) is a phase-retardation map of
the tissue, constructed form the two vectors S
1
and
S
2
. First, the average states S
1
and S
2
ref lected from
the glass–tissue interface were calculated, which de-
fined our input states. Next, a single rotation matrix
R was calculated that simultaneously transformed
S
1
Fig. 3. PS OCT images of in vivo human finger skin;
2 mm 3 2 mm, pixel size 5 mm 3 5 mm. Left to right,
Stokes parameters I, Q, U , and V for a single input po-
larization state. The images are gray-scale coded over a
40-dB range for I and from 1 to 21 for Q, U , and V.
September 15, 2000 / Vol. 25, No. 18 / OPTICS LETTERS 1357
Fig. 4. PS OCT images of in vivo human forearm skin,
2 mm 3 3.1 mm, pixel size 5 mm 3 5 mm for A, the
intensity image and 10 mm 3 10 mm for B, the phase-
retardation image. The images are gray-scale coded
over 40 dB for intensity and from 0 (black) to p (white)
for the phase retardation. For a description of the data
processing, see the text.
to S
1
and S
2
to S
2
at each position in the image. An
exact solution will not always exist. We determined
the best solution for R by calculating the angles f
1
be-
tween 共R ?
S
1
兲 and S
1
and f
2
between 共R ? S
2
兲 and S
2
and minimizing the sum of the angles, weighted by the
absolute length of the Stokes vector, 共f
1
I
1
1f
2
I
2
兲 j
min
.
The phase retardation, gray-scale coded from 0 to p,
is given by the rotation angle of matrix R. Although
the exact polarization states with which the sample is
probed are unknown, an image showing phase retar-
dation can be constructed from the measurements with
four different input states. The phase-retardation
image reveals birefringent structures at a depth of
300 mm below the surface, which we attribute to the
presence of collagen in the dermis of human skin.
Currently, a single-rotation matrix is calculated at
each depth in the sample; this assumes that the optical
axis is constant. In a more advanced approach, the
rotation matrix could be calculated between consecu-
tive Stokes vectors along a depth profile, which would
take into account variations in the orientation of the
optic axis with depth. The total encountered birefrin-
gence would be the sum of the absolute values of the
consecutive rotation angles. However, the presence of
speckle noise would likely lead to a significant overes-
timation of the total phase retardation.
In conclusion, we have demonstrated a f iber-based
PS OCT system that is capable of measuring birefrin-
gence of in vivo human skin. We achieved insensitiv-
ity to DPD in the sample arm fiber and orientation
of the optic axis in tissue by modulating across four
different input polarization states. Noninvasive mea-
surement of collagen birefringence could be an impor-
tant parameter in the assessment of burn depth, as
collagen is denatured as a result of thermal damage,
leading to a reduction in birefringence.
11
Research grants from the Whitaker Foundation
(grant 26083 to J. F. de Boer), the National Eye
Institute (1 R24 EY 12877-01), the U.S. Office of
Naval Research (N00014-94-1-0874), the Institute of
Arthritis, Musculoskeletal and Skin Diseases, the
U.S. Department of Energy, and the Beckman Laser
Institute Endowment are gratefully acknowledged.
J. F. de Boer’s e-mail address is deboer@bli.uci.edu.
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