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Exploring ultrasound-modulated
optical tomography at clinically useful
depths using the photorefractive
effect
Puxiang Lai, Yuta Suzuki, Xiao Xu, Lihong V. Wang
Puxiang Lai, Yuta Suzuki, Xiao Xu, Lihong V. Wang, "Exploring ultrasound-
modulated optical tomography at clinically useful depths using the
photorefractive effect," Proc. SPIE 8581, Photons Plus Ultrasound: Imaging
and Sensing 2013, 85812X (4 March 2013); doi: 10.1117/12.2003270
Event: SPIE BiOS, 2013, San Francisco, California, United States
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Exploring Ultrasound
-Modulated Optical Tomography at Clinically
Useful Depths Using the Photorefractive Effect
Puxiang Lai,
Yuta Suzuki, Xiao Xu
, and Lihong V. Wang
Washington University in St. Louis
Department of Biomedical Engineering, Optical Imaging
Laboratory
Campus Box 1097, 1 Brookings Drive, St. Louis,
Missouri 63130, USA
ABSTRACT
For years, ultrasound
-modulated optical tomography (UOT) has been proposed to image optical contrasts deep inside
turbid media (such as biological tissue) at an ultras
onic spatial resolution. The reported imaging depth so far, however,
has been limited, preventing this technique from finding broader applications. In this work, we present our latest
experimental explorations that push UOT to clinically useful imaging dep
ths, achieved through optimizing from different
aspects. One improvement is the use of a large aperture fiber bundle, which more effectively collects the diffused light,
including both ultrasound
-modulated and unmodulated portions, from the turbid sample a
nd then sends it to the
photorefractive material. Another endeavor is employment of a large aperture photorefractive polymer film for
demodulating the ultrasound
-induced phase modulation. Compared with most UOT detection schemes, the polymer film
based set
up provides a much higher etendue as well as photorefractive two
-beam
-coupling gain. Experimentally, we have
demonstrated enhanced sensitivity and have imaged through tissue
-mimicking samples up to 9.4 cm thick at the
ultrasonically-
determined spatial reso
lution
s.
Keywords
:
optical imaging,
ultrasound modulation,
ultrasound
-modulated optical tomography, acousto
-optic imaging,
photorefractive crystal,
photorefractive polymer
, tissue
-mimicking phantom, biological tissue, optical fiber bundle
１
1.
INTRODUCTION
As a non-
invasive, non
-ionizing, and relatively cost
-effective technique that is able to provide
functional information and
to distinguish
different
tissue types
, optical imaging
plays
a more and more important role in
bio
medical diagnosis
[1].
Its wide applications, however, are fundamentally
hindered
by the fact that light is highly diffused in biol
ogical tissue,
causing a trade
-off between the imaging depth and the resultant spatial resolution.
To overcome this limitation of reduced
spatial resolution due to strong light diffusivity when using optical imaging alone, researchers have proposed differe
nt
ways
to combine optical sensing with another imaging
modality
, e.g., ultrasound
, taking advantage of
the
contrast
provided by the pure optics and the spatial resolution
of the
ultrasound
[2]. Ultrasound
-modulated optical tomography
(UOT), also called acousto
-optic imaging, is such a
n example
[3,
4].
Further author information:
Puxiang Lai,
plai@wustl.edu
, 1 -314
-935 9587
Lihong V. Wang,
lhwang@wustl.edu
, 1 -314
- 935 6152
Photons Plus Ultrasound: Imaging and Sensing 2013, edited by Alexander A. Oraevsky, Lihong V. Wang,
Proc. of SPIE Vol. 8581, 85812X · © 2013 SPIE · CCC code: 1605-7422/13/$18 · doi: 10.1117/12.2003270
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Laser
(532 nm)
Variable
=
beam splitter
=
7
N
Mirror\
PR
material
Reference beam
Signal beam
0
is
QA\
r
' Beam block
High voltage
electric field
3.5 MHz,
10 -cycle burst
2.6 MPa (focal,p -p)
Ultrasound
transducer
In this technique, a focused
ultrasound field is applied to modulate the propagation of light inside
a turbid medium.
Photons are multiply diffused,
which leads to the formation of
random speckle
s outside the
sample
. Photons traversing
the ultrasound focal region
, however, experience
an
additional frequency shift that is equal to the
ultraso
und
frequency
,
through the displacements of scatterers and the variation in
the
medium
’s
index of refraction
[5, 6]. Since such
modulation only occurs
within the acousto
-op
tic interaction volume, and the depth of the modulation
is intrinsically
related
to local optical properties, detecting ultrasound
-modulated photons yields optically relevant information at the
ultrasonic resolution. Thus far, UOT has seen many application
s, such as conventional optical contrast imaging
[7],
multi
-wavelength functional imaging
[8], mechanical contrast imaging
[9, 10], fluorescence imagin
g [11], quantitative
measurement of optical properties
[12
], real
-time monitoring of thermal necrosis
[13
],
etc
. Most of these studies,
however, were performed in tissue
-mimicking phantoms or ex
-vivo tissue samples with
a limited depth less t
han the
clinically useful thickness
(usually
5-10 cm
[14
]). The depth insufficiency is
partially
caused by the fact that the number
of modulated photons is much less compared with the number of un
-modulated phot
ons, posing a s
trict
requirement to
extract
a very
small amount (typically 0.1
-1% of the total power) of MHz
-ordered frequency
-shifted modulated photons
from a much stronger untagged background that is
on the order of
THz
[15]. Moreover, s
ince photons travel
along
different optical paths inside the turbid medi
um, speckles form
ing
on the sample surface have random phase differences.
Therefore,
looking at one or only a few speckles results in
a low
flux of
tagged
-light. W
hile
directly detecting multiple
speckles onto a single
-element detector
does not coherently sum
up
the amount of tagged photons
, it
apparently
increase
s
useless untagged background
level
s. Different methods have been proposed to overcome this dilemma
and e
nhance the
signal detection sensitivity.
These
efforts include,
but are
not limited to, CCD
-based parallel detection
[16
, 17
],
photoref
ractive crystal (PRC)
-based interferometry
[18
, 19
], Fabry
-Perot interferometry
[20
, 21], as well as spectral hole
burning
-based filtering methods
[22
, 23]. In this work, we present our latest
explorations
[24
, 25] to p
ush UOT to
clinically useful thicknesses by
improving the etendue of diffused light collection and the efficiency of de
-modulation of
a photorefractive interferometric UOT system.
2.
METHO
DOLOGY
AND MATERIALS
Fig. 1 shows the experimental setup used in the
study. The detailed descriptions of this setup can be
found in
literatures
[24
, 25
], and thus are not iterated in thi
s report. Several things, however, need to be specially
described
.
Fig.
1. System
experimental
setup
.
Y
is the optical illumination direction, and
Z
is the ultrasound propagation direction
.
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core
interfiber
cladding
space
First,
the diffused
light
from the experimental sample
, includ
ing both
ultrasound
-modulated and unmodulated portions,
was collected
using
an optical fiber bundle from
Edmund Optics (NT38
-659). A
s illustrat
ed in Fig. 2,
multiple optical
fibers were tightly packed into the round bundle,
which had
a net clear aperture o
f 12.7 mm, a length of 305 mm
, a high
NA of 0.55, and a full acceptance angle of 68
o
. This fiber bundle offered
improved light throughput (or etendue)
compared with standard single fiber assemblies, and greater flexibility
/convenience
compared with regular
lenses
used in
free space
.
Fig.
2 Illustration of the optical fiber bundle used in this study
Second,
two types of photorefractive materials
, as shown in Fig. 3,
were used in this study to efficiently convert the
ultrasound
-induced
frequency and
phas
e modulation into
an
intensity modulation that could be detected by a photodiode
[18
, 19, 26]. T
he first type was a Bi
12
SiO
20
(BSO)
photorefractive
crystal
from Elan, Russia
that
had a dimension of
10×10×5 mm
3
along
the
X
,
Z
,
Y
directions
. Under our experimental conditions (described later), the crystal yielded
a
response speed of ~100 ms, and two
-wave
-mixing (TWM) gain coefficients of 0.81 and 0.26 cm
-1
, respectively, with and
without a 2.1 kHz, 8 kV/cm high voltage AC electronic field
applied across the crystal (along the
X
direction)
. The
second type is a phot
orefractive polymer (PRP) film from Nitto Denko Technical, CA, USA.
The
PRP
, 50.8×50.8×0.1
mm
3
along
X
,
Z
, and
Y
, respectively,
had
much larger dimensions along the
X
and
Z
directions
than the BSO
, as shown in
Fig. 3.
In our study, a 400
-1000 kV/cm high vo
ltage DC electronic field was applied
along the PRP
(along the
Y
direction)
to enhance the TWM gain coefficient up to 9.2 cm
-1
. However,
it must be noted that this polymer film was relatively slow,
with a response time up to several seconds under our exper
imental conditions, which is not desirable for future
in-vivo
applications to compensate the speckle decorrelation caused by
the
physiological motions.
However
, to demonstrate
deep
imaging in a non
in-vivo
scenario, this slow
response speed was practical
ly adequate, as will be shown later in this
report.
Fig. 3 Dimension comparison of the BSO photorefractive crystal and the photorefractive polymer (PRP) film
At this point, it is necessary to highlight the advantage of using this large aperture PRP in terms of diffused light
collection etendue. Etendue is an important parameter used to specify the geometric capability of an optical system to
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100001
1i
0.11
n
PRCs
rl
Ii
h
i
e.o ,tJ
5
.0.\
aJ eto
Q
,e
\V
O
Q
<z",
a
..Qo
Q
'\
c.)\
Qea ot
e;o
h
r
transmit, and/or receive radiation (in our case, light), its throughput. Numerically, it is defined as
2
sin ( / 2)
GA
,
where
A
is the active surface area of the optical component, and
Ω
represents the emission/acceptance angle. In UOT, the
experimental samples usually have apertures or diameters of several centimeters, and accordingly large output etendue
s
of 5,000-30,000 mm
2
sr, considering an Ω of 180
o
as light gets multiply scattered within the experimental sample. As
shown in Fig. 4, previous detections schemes that allowed coherent signal summation over multiple speckles
, including
confocal Fabry-Perot interferometers (CFPI), PRC-based interferometers
, and spectral hole-burning crystals, had much
smaller etendues due to their limited dimensions and/or acceptance angles. It meant only a small portion of tagged
photons were effectively used and contributed to the final UOT signal levels. This, accompanied with the factors
discussed earlier in the Introduction, makes it challenging to image through thick turbid media within the illumination
safety limits for sound [
27
] and light [
28
]. Comparably, the PRP film used in this study provided an etendue up to
400
-500 mm
2
sr, imperfect yet much larger than previous schemes, promising a manifold increase in the UOT
signal-
to-noise ratio due to the increase of useful tagged photons.
Fig. 4 Comparison of etendue for experiment sample and different detection schemes. Each column represents one case reported in
literatures: confocal Fabry-Perot interferometers (CFPI) [
20, 21
], BSO crystals [
9, 18, 24], GaAs crystals [
12, 19, 29], SnPS crystals
[30], spectral hole-burning crystals [22
, 23], and the PRP film used in this work. The error bar of the experimental sample comes from
the variance of sample surface dimensions (ranging from 4 ×4 to 10×10 cm
2
), while the error bars for the PRCs are from the
approximation of Ω
(20-
40
o
according to Ref. [23
]).
In this study, two phantoms that mimicked optical properties of tissue were used as the experimental samples. These
phantoms were 6 cm and 9.4 cm thick, respectively. Both were composed of 10% (by weight) porcine gelatin
(Sigma-Aldrich, MO, USA), 89% water, and 1% Intralipid (diluted from 20% Intralipid from Fresenius Kabi, Germany)
.
The resultant
'1
10
s
cm
, measured by oblique-incidence reflectometry [
31
], and the absorption coefficient μ
a
was
0.12 cm
-1
, measured by a spectrophotometer (Cary 50, Varian, CA, USA)
. To
provide optical contrasts inside these two
phantoms, absorbing objects were embedded at the central plane of the samples. The absorbers were made from the same
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1
1
1
1
1
20 40 60
80
Z (mm) = Time (ps) x 1.5 mm/ps
material with the background but dyed with India ink to have a higher absorption coefficient at 2 cm
-1
. More details
about these samples and how the absorbers were deployed will be given later.
Fig. 5 shows two examples of normalized UOT signal waveform when the ultrasound burst from a 3.5 MHz transducer
(Panametrics A318S, Olympus NDT, MA, USA) propagated outside and across an optical absorber that was embedded
inside the 6-cm thick phantom. As seen, when the ultrasound burst travelled away from the absorber (the green curve),
the modulated-light power increased with the
in situ
pressure amplitude, forming a summit around 38-
39
μs when the
burst arrived at the ultrasound focal position (38.1 mm away from the transducer surface). With the presence of the
absorber, however, the modulated-light power began to reduce when the ultrasound burst encountered the absorbing
region, due to the decreased amount of modulated photons and the efficiency of ultrasound modulation [
6, 12, 32
].
Therefore, the valley of the UOT signal is seen as the signature of the optical absorber.
Fig. 5 Examples of UOT signal waveform with ultrasound pulses propagating outside (green curve) and across (red curve) an optical
absorber embedded inside the 6-cm thick phantom.
3.
RESULTS AND DISCUSSION
Fig. 6(a) shows the cross
-section
photograph
of the 6
-cm thick tissue
-mimicking phantom mentioned earlier. There
were
three
absorbing object
s that had
a higher optical absorption coefficient (
2 cm
-1
) th
an the background
. This cross
-section
was embedded
at
Y
= 3 c
m
in the 6
-cm thick sample.
Before the experiment, the ultrasound focus was aligned to intersect
with the needle on the left in both
Y
and
Z
directions. During the experiment,
the BSO
photorefractive
crystal was used
in the detection scheme
[24
]. And
a signal
beam of 35 mm diameter, 208 mW/cm
2
, a reference beam of 10 mm diameter,
30 mW/cm
2
, and a 10
-cycle burst with a 2.6 MPa (peak
-to-peak) focal pressure
, and a repetition rate of 100 Hz,
were
employed.
Two groups of measurements were performed. First, the
lig
ht and sound
components
remained stationary,
but
the sample scanned along the
X
direction
at a step size of
0.3175 mm.
In the second group of measurements, the light and
the sample kept still, but the ultrasound focus was scanned along the
Y
direction with
the sample fixed at the position of
X
= 6.35 mm.
At each position, one UOT scan line (“A line”) as shown in Fig. 5 was obtained
, averaging over 64
ultrasound bursts
. Fig. 6(b) and (c) show t
he resulting 2
-D UOT image
s, with each scan line normalized to it
s peak
power for better visualization contrast
. Note the distance along the
Z
direction was converted from the product of
temporal position with the sound speed in the phantom (~1.5 mm/μs)
. In Fig. 6(b), the three embedded absorbers can be
clearly seen in
the
XZ
plane, even though Obj3 was only partially imaged due to the limited scanning range of our
translation stages. In Fig. 6(c), Obj1 was also clearly revealed in the
YZ
plane. Moreover,
the
profiles (Fig. 6 d
-f) along
the
dashed lines in Fig. 6(b) and
(c)
provided
more information,
like
the dimensions of the objects, the interspacing
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(a)
(b) 50
(e)
45-
E
Z 40-
Needle
Objl
Obj2 Obj3
I
n
35
0 5
10
15
20
25
X(mm)
5
10
15
20 25
X (mm)
3 1.0
0
a
1
- 0.8
0.6
0.4
(c) 50
45-
E
7; 40-
35
Ah.
Obj1
(f)
1
0.8
0.6
0.4
0
0.4
50
45 40
Z(mm)
35
4 8
12
Y (mm)
(f)
1.1
r 0.9
v
0.7
v 05
0 0
4
8
Y (mm)
12
between the objects, and the resolutions along the
X
and
Z
directions.
For example, in Fig. 6(d), Obj1 was centered at
X
≈
6.0 mm, ~2.4 mm wide
(all imaged sizes in this st
udy were determined from the full
-width at half maximum of the
contrasts)
, and
Obj2 at
X
≈ 16.9 mm, ~4.3 mm wide.
These were quite
consistent with
what can be seen from the
cross
-section photograph
(Fig. 6a)
that Ob1 was ~2.1 mm, Obj2 was ~4.0 mm, and thei
r center
-to-center distance was
~10 mm.
The resolution
of UO
T along the
X
direction, estimated from
the 1/4
-3/4
of the imaging contrasts, was
~1.05
mm, which was only slightly wider than the ultrasound focal width (0.875 mm).
In Fig. 6(e), Obj1 was about 5
.1 mm
long,
and was
, again
, reasonably consistent with its actual length of 4.5 mm
shown in
the photograph. The imaging
resolution along the
Z
direction was ~2.4 mm, which is 0.5
-0.6 of the ultrasound burst length (10 cycles of 3.5 MHz is
about 4.3 mm long
spatially).
Lastly, Obj1 was ~3.8 mm wide from the profile shown in Fig. 6(f), a value very close to
its actual dimension (3.5 mm). The estimated imaging resolution in the
Y
dir
ection was ~1.17 mm, and was close to
the
focal width of the ultrasound field.
Fig.
6 (a) The cross
-section
photograph of the 6
-cm thick phantom.
(b ) The resultant 2
-D
UOT
image of the phantom in the
XZ
plane
with
Y
= 3 cm.
(c) The 2
-D UOT image in a
YZ
plane across Obj1 (
X
= 6
.35
mm in Fig. 6b)
. (d
-f) Modulated light power distr
ibutions
along dashed lines labeled in Figs. (b
) and (
c), respectively.
In (d) and (f), the blue squares are the measured data, and the red curves
the FFT smoothed results.
Reproduced from Ref.
[24].
To
image
through the 9.4
-cm thick phantom sample (its cross
-section shown in Fig. 7a), the photorefractive polymer film
discussed earlier was used to enhance the detection sensitivi
ty without
the need of em
ploying
much more intense light
and sound illuminations
[25
]. Except for the
replac
ement
of the photorefractive material, other experiment parameters
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(a)
remained
very similar or unchanged from
the last
measurements performed for Fig. 6. For example, the used signal beam
had a diameter of 24 mm, an intensity of 192 mW/cm
2
, and the reference beam a diameter of 30 mm, an intensity of 9.9
mW/cm
2
. The ultrasound field,
the
averaging
times
of UOT signal, the way of scanning the experimental sample, and the
signal processing methods were
kept
identical to those used for Fig. 6. The resulting “B
-
Mode” 2
-D UOT image of the
cross
-section was shown in Fig. 7(b), from which one
can tell the existence, dimensions, inter
-spacing of the two
embedded absorbers
, and
the resolutions in the
XZ
plane. Note that this UOT image was obtained with a relatively
straightforward and
inexpensive
apparatus in a tissue
-mimicking phantom sample wi
th a thickness of 9.4 cm,
which is
equivalent to
94 transport mean free paths
, or 9.4 cm thick
ness of
normal human breast tissue, a clinically useful depth
that had never been achieved, not
to mention
explored, until most recently
[23
-25
].
Fig.
7 (a)
Cross
-section of
the 9.4 cm
-thick phantom sample, embedded with two absorbing objects at its central
Y
plane (
Y
= 4.7 cm)
.
(b) The resulting
2-D UOT image of
the
central
cross
-section
. Reproduced from Ref.
[25].
4.
SUMMARY AND FUTURE WORK
In summary, we
reported our latest improvements to a relatively simple and cost
-effective p
hotorefractive
-bas
ed UOT
system
. Our efforts included the use of
a large aperture optical fiber bundle to improve the diffused light collection
convenience and efficiency, and a photorefractive polymer to improve the etendue and two
-wave
-mixing gain coefficient
for
coherent
de
-modulation of
ultrasound
-induced frequency/phase modulation. E
xperimentally
, we demonstrated UOT
imaging at clinically useful thicknesses up to 9.4 cm in tissue
-mimicking phantom samples
, within
the
light
[28
] and
sound
[27]
safety limits.
That being said,
some
more
work need
s to done to make
our apparatus
towards clinical
applications.
For example,
the syst
em need
s to be modified to work at a more tissue
-friendly optical wavelength such as
700
-800 nm or 1064 nm for deeper
penetration
depth in tissue,
benefitting from less optical tissue attenuation and higher
ANSI safety limits
[28
]. Moreover, faster responding photorefractive materials
(e.g. GaAs
crystal
[33
], SnPS
crystal
[30
],
and fast polymer
s [34
]) are desired to
better
compensate the
in vivo
speckle decorrelation (1
-10 kHz) from physiological
mot
ions.
ACKNOWLEDGEMENT
S
The authors
thank Sandra Matteucci for editing the manuscript, and
Nitto Denko Technical at Oceanside, CA for
providing the photorefractive polymer film used in
the current
study.
This research is sponsored in part by the Nationa
l
Academies Keck Futures Initiative grant IS 13
, Natio
nal Institute of Health
grants
DP1 EB016986 (NIH Director’s
Pioneer Award),
R01 EB000712 and U54 CA136398. L.W. has a financial interest in Microphotoacoustics, Inc. and
Endra, Inc., which, however, did
not support this work.
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