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Two-dimensional birefringence imaging in biological tissue by polarization-
sensitive optical coherence tomography
de. Boer, J.F.; Milner, T.E.; van Gemert, M.J.C.; Nelson, J.S.
DOI
10.1364/OL.22.000934
Publication date
1997
Published in
Optics Letters
Link to publication
Citation for published version (APA):
de. Boer, J. F., Milner, T. E., van Gemert, M. J. C., & Nelson, J. S. (1997). Two-dimensional
birefringence imaging in biological tissue by polarization-sensitive optical coherence
tomography.
Optics Letters
,
22
(12), 934-936. https://doi.org/10.1364/OL.22.000934
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934 OPTICS LETTERS / Vol. 22, No. 12 / June 15, 1997
Two-dimensional birefringence imaging in biological tissue by
polarization-sensitive optical coherence tomography
Johannes F. de Boer
Laser Center, Academical Medical Center, Meibergdreef 9, 1105 AZ Amsterdam, The Netherlands, and
Beckman Laser Institute and Medical Clinic, University of California, Irvine, Irvine, California 92612
Thomas E. Milner
Beckman Laser Institute and Medical Clinic, University of California, Irvine, Irvine, California 92612
Martin J. C. van Gemert
Laser Center, Academical Medical Center, Meibergdreef 9, 1105 AZ Amsterdam, The Netherlands, and
Beckman Laser Institute and Medical Clinic, University of California, Irvine, Irvine, California 92612
J. Stuart Nelson
Beckman Laser Institute and Medical Clinic, University of California, Irvine, Irvine, California 92612
Received January 13, 1997
Using a low-coherence Michelson interferometer, we measure two-dimensional images of optical birefringence
in bovine tendon as a function of depth. Polarization-sensitive detection of the signal formed by interference
of backscattered light from the sample and a mirror in the reference arm give the optical phase delay between
light that is propagating along the fast and slow axes of the birefringent tendon. Images showing the change
in birefringence in response to laser irradiation are presented. The technique permits rapid noncontact
investigation of tissue structural properties through two-dimensional imaging of birefringence. 1997 Optical
Society of America
The demand for noninvasive optical imaging in bio-
logical tissue has led to the development of several
techniques to circumvent the common problem of
scattering in turbid media; such techniques include
diffusing-wave spectroscopy,
1
time-gated imaging
using snakelike photons,
2
two-photon fluorescence
imaging,
3
ultrasonic modulation of diffusing waves,
4,5
and optical coherence tomography
6,7
(OCT).
OCT uses the partial coherence properties of a
light source to image structures with high resolution
(1–15 mm) in turbid media such as biological tissue.
The sample is positioned in one arm (the sample
arm) of a two-beam interferometer. Interference
fringes are formed when the optical path length
of light backscattered from the sample matches
that from the reference to within the coherence
length of the source light. The optical path length
in the reference arm acts as a gate on the detec-
tion, selecting only light backscattered from the
sample that has traveled the same optical path
length. By lateral and longitudinal scanning, two-
dimensional (2D) OCT images are constructed that
map the amplitude of light backscattered from the
sample. Lateral resolution is determined by the spot
size of the beam focus of incoming light and the collec-
tion aperture; longitudinal resolution is determined by
the coherence length of source light.
In this Letter we present a combination of OCT
and polarization-sensitive detection
8
to record 2D
images of the change in polarization of circularly
polarized incoming light backscattered from a turbid
birefringent sample. In contrast with conventional
OCT, in which the magnitude of backscattered light as
a function of depth is imaged, we use the backscattered
light to image the magnitude of the sample birefrin-
gence as a function of depth. 2D maps of birefringence
of biological materials can reveal important structural
information that is difficult to resolve with other imag-
ing techniques. Partial loss of birefringence is known
to be an early indication of tissue thermal damage
9
(e.g., burns or laser treatments). To demonstrate
polarization-sensitive OCT, we present 1-mm-wide by
700-mm-deep images of bovine tendon birefringence
before and after pulsed laser irradiation.
Figure 1 shows a schematic of the polarization-
sensitive OCT system used in our experiments. Light
passes through a Glan–Thompson polarizer to se-
lect a pure linear vertical input state and is split
into reference and sample arms by a polarization-
insensitive beam splitter (ref lection and transmission
coefficients for linear vertical and horizontal polariza-
tion states were 0.5 6 0.05). Light in the reference
arm passes through a zero-order quarter-wave plate
(QWP), rotating at
Ù
Q 200p radys. Following reflec-
tion from a planar mirror and a return pass through
the QWP, light in the reference arm has a rotat-
ing linear polarization (400p radys). For improved
signal–noise ratio,
10
a neutral-density filter positioned
in the reference arm reduces intensity noise by a fac-
tor of 50. Light in the sample arm passes through
0146-9592/97/120934-03$10.00/0 1997 Optical Society of America
June 15, 1997 / Vol. 22, No. 12 / OPTICS LETTERS 935
Fig. 1. Schematic of the polarization-sensitive OCT setup.
SLD, superluminescent diode, 0.8-mW output power, cen-
tral wavelength l
0
862 nm, spectral FWHM Dl 21 nm
s l
0
2
V
p
ln 2ypcd; P, polarizer; BS, beam splitter; QWP’s,
quarter-wave plates; NDF, neutral-density filter; L, lens.
2D images are formed by longitudinal movement of the
sample with constant velocity v 1 mmys(zdirection), re-
peated after each 10-mm lateral displacement (x direction).
a QWP oriented at 45
±
to the incident vertical polar-
ization to give circularly polarized light. After double
passage through a lens and the sample, and propaga-
tion through the QWP, light in the sample arm is in an
arbitrary (elliptical) polarization state, determined by
the sample birefringence.
The intensity of the light incident upon the detector
is given by recombination of the light in both arms of
the interferometer:
kIl kI
r
l 1 kI
s
l 1 2 RefkC
p
r
st, z
r
d ? C
s
st, z
s
dlg , (1)
where Cst, zd is the light-amplitude field vector,
subscripts r and s denote the reference and the sample
arms, respectively, z
r, s
are the optical path lengths,
and the angle brackets denote ensemble averaging.
The source spectral density Ssvd is assumed to be
Gaussian h~ expf2sv2v
0
d
2
yV
2
gj with FWHM
2V
p
ln 2radys. The evolution of the polarization
state in each arm of the interferometer is computed
with the Jones matrix formalism, where we neglect
optical dispersion in the sample and spectral depen-
dence of the zero-order QWP’s over the spectrum of the
source. The interference intensity between light in
the sample and the reference arms can be separated
into horizontal, A
H
, and vertical, A
V
, components
that are proportional to the light-amplitude fields
backscattered from the sample:
RefkC
p
r
st, z
r
d ? C
s
st, z
s
dlg A
H
1 A
V
,
A
H
ø sins2
Ù
Qtdcoss2k
0
Dzdexpf2sVDzycd
2
gcossk
0
zdd ,
A
V
ø coss2
Ù
Qtdcoss2k
0
Dz 1 2ad
3 expf2sVDzycd
2
gsinsk
0
zdd, (2)
where Dz is the optical path-length difference between
the sample and the reference arms of the interferome-
ter. In our system Dz is a function of time and longitu-
dinal velocity of the sample, z is the backscatter depth
in the sample, k
0
2pyl
0
, c is the speed of light in
vacuum, d is the difference in refractive indices along
the fast and the slow axes of the birefringent sample
sn
s
2 n
f
d,
Ù
Q is the rotation speed of the QWP in the ref-
erence arm, and a is the angle of the fast optical axis of
the birefringent sample with the horizontal. The only
approximation in Eqs. (2) is that the product of bire-
fringence and propagation depth in the sample (in the
present case zd#2.6 mm) is much smaller than the
width of the coherence envelope s2cyV 19 mmd.
Phase-sensitive demodulation of the recorded signal
with respect to the angular position of the rotating
QWP s
Ù
Qtd permits separation of the intensities corre-
sponding to vertically and horizontally polarized light.
In addition to the carrier frequency fcoss2k
0
Dzdg within
the coherence envelope hexpf2sVDzycd
2
gj both signals
show an oscillation with a periodicity determined by
the product of sample birefringence sdd and propaga-
tion depth szd that allows for birefringence imaging.
We made scans by moving the sample at constant
velocity v 1mmys, giving a carrier frequency n
2vyl
0
2.4 kHz. To form 2D images, we recorded a
longitudinal scan after each 10-mm lateral displace-
ment of the sample. The 4-mm-diameter beam, fo-
cused upon the sample by a lens s f 50.2 mmd,gavea
14-mm beam-waist diameter. In air, the sample arm
was matched in length to the reference arm at a posi-
tion 200 mm past the focal point, leading to a matched
length in the beam focus ,400 mm deep in a sample
with refractive index n 1.4.
11,12
The detector was ac
coupled, and the signal was amplified, high-pass fil-
tered at 1 kHz with 18-dByoctave roll-off, and digitized
with 16-bit resolution at 50,000 points per second.
Signal processing consisted of squaring the detected
signal and phase-sensitive demodulation with respect
to
Ù
Qt to separate the horizontal, I
H
, and vertical, I
V
,
components of the backscattered light. Then, within
each longitudinal scan, data points were averaged
with a Gaussian weight function (FWHM 700 data
points) to form one image pixel. The resulting sig-
nals give the backscattered horizontal and vertical in-
tensities as a function of depth z with a resolution of
10-mm physical distance, modulated with their respec-
tive birefringence-dependent terms:
I
H
szd ~ cos
2
sk
0
zdd, I
V
szd ~ sin
2
sk
0
zdd . (3)
In Fig. 2a, a birefringence image of fresh bovine
tendon is shown. Measurements on 1cm32cm
samples at least 1 cm thick were done within 48 h
of sacrifice of the bovine. The banded structure,
indicative of birefringence, is clearly visible up to
a physical depth of 700 mm. By measuring the
optical-versus-physical thickness of a thin slice,
11
we
found the average refractive index of the tendon to be
¯
n 1.42 6 0.03. We determined the birefringence
by measuring the average distance between the start
(zero crossing) of the first and the second yellow
bands from the top of the figure over the full width
of Fig. 2a (100 lateral scans). The average distance
¯
z 116 6 13 mm corresponds to a full period of
the sine squared in relations (3), k
0
¯
zd p. The
measured birefringence d 3.7 6 0.4 3 10
23
of
bovine tendon (predominantly type I collagen) is in
agreement with reported values of 3.0 6 0.6 3 10
23
(Ref. 13) and 2.8–3.0 3 10
23
.
14,15
Fitting exps22zygd,
between z 150 2 600 mm depth, to the total
backscattered intensity in the sample, I
b
szd ~
I
H
szd 1 I
V
szd, averaged over the image in Fig. 2a
936 OPTICS LETTERS / Vol. 22, No. 12 / June 15, 1997
Fig. 2. Images of fresh bovine tendon 1 mm wide by 700 mm deep. Each pixel represents a 10 mm 3 10 mm area.
The dynamic range of the system was 48 dB. We generated false color birefringence images a and b by computing
sgnfI
H
szd 2 I
V
szdg10 logjI
H
szd 2 I
V
szdj. Noise is represented by deep blue or deep red. The color scale at the right gives
the magnitude of signals. The banded structure, indicative of the birefringence, is clearly visible. a, Birefringence
image of fresh bovine tendon. b, Birefringence image of bovine tendon following exposure to three consecutive 1-J
150-ms laser pulses (l 1.32 mm) spaced by 10 ms, incident from the upper left at 35
±
with respect to the surface
normal. The beam diameter was 2 mm. Initial surface temperature after laser irradiation was 77
±
C, dropping to 61
±
C
after 0.25 s. c, Total backscattered intensity image of bovine tendon, constructed from the same measurement shown
in b to demonstrate differences from the polarization-sensitive image. We generated a false color image by computing
10 logfI
H
szd 1 I
V
szdg. The color scale to the right gives the magnitude of the signals in c.
(100 lateral scans), gives g ø 0.2 mm. Decay of the
total backscattered light intensity with depth depends
on several factors, among them attenuation of the
coherent beam by scattering and the geometry of the
collection optics.
In Fig. 2b, a birefringence image of laser-irradiated
bovine tendon is presented. The image clearly shows
a decrease in the birefringence at the center of the
irradiation zone, extending into the tendon over the
full depth of the image (700 mm). Further, the direc-
tion of incoming laser light (from the upper-left cor-
ner, at an angle of 35
±
with the normal of the surface)
is observed. The surface temperature of the tendon
was monitored during laser irradiation by IR radiome-
try. For comparison, Fig. 2c shows an OCT image of
the total backscattered intensity I
b
szd. Although less
backscattered light from the irradiated area can be ob-
served, the polarization-sensitive image (Fig. 2b) re-
veals important structural information not evident in
Fig. 2c.
We have shown that polarization-sensitive optical co-
herence tomography can reveal structural information
on birefringent turbid media such as biological tissue
that is not available when polarization-insensitive OCT
is used. Polarization-sensitive OCT has the potential
to provide guidance regarding optimal dosimetry for
thermally mediated laser therapeutic procedures by
permitting real-time diagnostics at each irradiated site
through detection of changes in birefringence associ-
ated with thermal damage and pathological conditions.
This would permit a semiquantitative evaluation of the
efficacy of laser therapy as a function of incident light
dosage.
Research grants from the Biomedical Research Tech-
nology Program and the Institute of Arthritis and
Musculoskeletal and Skin Diseases of the National
Institutes of Health, the Whitaker Foundation (WF
21025), and the Dermatology Foundation are grate-
fully acknowledged, as is institute support from the
U.S. Department of Energy, the National Institutes
of Health, and the Beckman Laser Institute Endow-
ment. In addition, J. F. de Boer (JFdB) and M. J. C.
van Gemert gratefully acknowledge financial support
from the Dutch Technology Foundation (STW, grants
AGN 33.2954 and AGN 55.3906) and the Academical
Medical Center. JFdB thanks R. Sprik for stimulat-
ing discussions.
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