Video-rate three-dimensional optical coherence
tomography
Markus Laubscher, Mathieu Ducros*, Boris Karamata, Theo Lasser and René Salathé
Institute of Applied Optics, Swiss Federal Institute of Technology, CH-1015 Lausanne, Switzerland
Markus.Laubscher@epfl.ch
* now with the Ophthalmology Department, The University of British Columbia,
Vancouver British Columbia, Canada, V5Z 3N9
Abstract: Most current optical coherence tomography systems provide
two-dimensional cross-sectional or en face images. Successive adjacent
images have to be acquired to reconstruct three-dimensional objects, which
can be time consuming. Here we demonstrate three-dimensional optical
coherence tomography (3D OCT) at video rate. A 58 by 58 smart-pixel
detector array was employed. A sample volume of 210x210x80 µm3
(corresponding to 58x58x58 voxels) was imaged at 25 Hz. The longitudinal
and transverse resolutions are 3 µm and 9 µm respectively. The sensitivity
of the system was 76 dB. Video rate 3D OCT is illustrated by movies of a
strand of hair undergoing fast thermal damage.
2002 Optical Society of America
OCIS codes: (170.4500) Optical coherence tomography; (170.3880) Medical and biological
imaging; (120.3890) Medical optics instrumentation; (170.6900) Three-dimensional
microscopy; (110.6880) Three-dimensional image acquisition
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1. Introduction
Over the past 15 years the biomedical imaging technique called optical coherence tomography
(OCT) has experienced many technological improvements and found a host of useful
applicationss [1-3]. The main aspects of development have been spatial resolution [4],
sensitivity and acquisition speed [5,6]. In addition, new implementations of OCT have been
developed that provide additional information about the sample under study. For example,
polarization-sensitive OCT allows the measurement of depth-resolved sample birefringence
[7,8] and Doppler OCT permits the assessment of flow velocity in biological samples [9].
However, acquiring two-dimensional OCT images is not sufficient to fully describe the three-
dimensional morphology of biological samples under study. For example, three-dimensional
OCT images would be required to measure the depth and lateral extent of epithelial tumors in
the skin, cervix or oral mucosa. A few research groups reported the reconstruction of three-
dimensional maps of reflectivity obtained by acquiring two-dimensional arrays of adjacent
OCT A-scans, but such a procedure can be time consuming[10-12].
Besides the “classic” longitudinal OCT imaging technique based on A-scans, two classes
of en face (transversal) OCT imaging techniques have been proposed: the “flying spot” and
the “parallel OCT” techniques. Both can be used to acquire three-dimensional reflectivity
maps. In the “flying spot” technique the probing light beam is scanned transversally in raster
scans to acquire en face images at different depths [13,14] whereas in the “parallel OCT”
technique wide-field illumination and acquisition is used [15-17]. The detectors employed in
the latter technique are photodetector arrays in contrast to all other OCT techniques which
rely on single-unit detectors. The need for lateral scanning is in this case eliminated, to the
advantage of higher acquisition rates.
Charge coupled device (CCD) cameras are the most commonly used imaging devices for
parallel imaging schemes. However, CCD cameras suffer from two drawbacks when used in
parallel OCT systems: (1) the high optical DC intensity reflected by the reference mirror
reduces the dynamic range available for AC interferometric signal detection, (2) the CCD
frame rate (typically ~100 Hz for 512x512 pixels) is much lower than the interferometric
signal frequency (typically greater than 1 kHz). In this case a lock-in detection or synchronous
illumination scheme has to be employed [15], which limits the image acquisition speed. A
different photodetector array based on CMOS technology was specifically developed for
parallel OCT [17,18]. Besides transducing light signals into electrical signals, CMOS
detectors offer the additional functionality of customized, integrated signal processing for
each pixel. Optical coherence tomography with a parallel detection scheme using such one-
and two-dimensional smart pixel detector arrays (SPDA) was previously demonstrated on
reflective surfaces [17]. Recently, we have shown the feasibility of using a SPDA in scattering
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samples as well [19]. In the present work we use a SPDA to demonstrate for the first time to
our knowledge 3D OCT imaging at video-rate.
2. Method
2.1 Optical set-up
The optical set-up is illustrated in Figure 1. The light source employed is a compact
femtosecond mode-locked Ti:Sapphire laser (MLTS) (FemtoLasers Inc., Vienna, Austria)
with a nearly Gaussian spectrum centered at 800 nm and a full-width-at-half-maximum
(FWHM) spectral bandwidth of 100 nm. Lenses L1 and L2 form a telescope to increase the
beam diameter before it enters a free space Michelson interferometer. An average power of
430 mW is available at the interferometer input. A beamsplitter cube (BS) separates the light
into the interferometer reference and sample arms. A variable neutral density filter wheel (F)
is placed into the reference arm and a compensation glass plate (C) of equal thickness into the
reference arm. Two identical microscope objectives (L4 and L5, 20x) are used to illuminate
and collect reflected light from the sample (S) and reference mirror (RM). The incident
average power on the sample is 120 mW. The illumination profile on the sample is
approximately gaussian and covers the square field of view of the detector of 210x210µm2.
Light reflected from S and RM interferes only if the optical path lengths match to within the
source coherence length. RM is translated longitudinally using a voice-coil scanning stage
(Physik Instrumente (PI) GmbH & Co) that is driven by a triangular input function at a
frequency of 12.5 Hz. The scan amplitude is 80µm, as measured by the voice coil stage
encoder. Acquisition is performed both during the forward and backward half-period of the
triangular scan, i.e. at 25 Hz. The sample is imaged by lens L6 onto a SPDA with 58x58
pixels. Each pixel consists of a silicon photodiode coupled to a CMOS electronic circuit that
amplifies and demodulates AC signals [17]. Each pixel output provides an analog voltage
proportional to the envelope of the optical interference signal. The analog signals
corresponding to each pixel are read out sequentially at a rate of 5 MHz, digitized by a 12-bit
A/D card and displayed on a computer screen. Volumetric datasets with 58x58x58 pixels,
corresponding to 210x210x80 µm3 are thus acquired at a rate of 25 Hz.
Fig. 1. Parallel OCT optical setup schematic. The different elements are: mode-locked
Ti:Sapphire femtosecond laser (MLTS); achromatic lenses (L1, L2, L3, and L6); non-
polarizing achromatic beamsplitter cube (BS); identical achromatic microscope objectives 20X,
0.45 NA (L4 and L5); reference mirror (RM); variable neutral density filter wheel (F);
compensation glass plate (C); 58 by 58 smart pixel detector array (SPDA) and sample (S).
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2.2. Sample
The sample was a strand of dark human hair on a microscope glass slide positioned at the
focal plane of the microscope objective L4. The average irradiance incident on the sample is
approximately 173 W/cm2. Because of the high melanin concentration in the hair strand the
illuminating laser beam is strongly absorbed. As the hair is only in contact with a glass plate
and with the air the heat transfer to the surrounding media is rather low and the accumulated
thermal energy causes the hair to swell and to burn. Only because of the high absorption
coefficient of the imaged sample do we observe an interaction between the probing laser beam
and the sample. Indeed, we have observed light-colored hair and onions under identical
conditions and no damage to the sample was observed. Even though the total power used for
sample illumination is high (120 mW) the irradiance remains relatively low. Indeed, the
employed average irradiance is inferior to irradiance reported in literature in point by point
scanning OCT systems using femtosecond lasers[4]. Nevertheless, thermal effects, depending
on irradiance, pulse length, exposition duration, tissue absorption coefficient and thermal
properties, should be investigated for each specific sample.
3. Results
3.1. System performance
We measured the system longitudinal response on a mirror in air to be 3 µm (FWHM), which
is in good agreement with the theoretically expected value for a Gaussian spectrum of 100 nm
bandwidth at a central wavelength of 800 nm. Transverse resolution in air was determined
using a USAF resolution target at the focal plane of L4. Using the 20x objectives we could
resolve reflective bars with a maximum spatial frequency of 114 mm-1, corresponding to a
transverse resolution of 8.8 µm. The transverse resolution is limited by the NA of the
microscope objectives (0.45) and the fill factor of the detector array (10%).
To measure the system sensitivity we imaged an air-water interface (2% power reflection)
and varied the attenuation of the reference arm intensity by rotating the neutral density filter
wheel F until the detector signal was maximized to a value that we call V2% max. The
sensitivity S in decibel of the system is then given by
2% max 120 log 10 log
0.02
V
S
σ
= + (1)
where σ is the electronic signal noise that is experimentally taken to be the standard deviation
of all voxels values when no sample was present. The sensitivity was measured to be 76 dB.
During all experiments on biological samples, such as the hair strand, the neutral density filter
wheel stayed at the same position.
The advantages of using the method described above to measure the system sensitivity are
threefold: (1) The sample (water) mimics typical biological samples reflectivities; (2) The
electronic AC gain of the SPDA detector depends on the level of optical DC illumination. The
lower the DC illumination the higher the AC gain. Therefore, by decreasing the reference arm
intensity the AC gain is increased more than the optical AC signal is attenuated and thus a
higher electric output signal can be obtained; (3) By decreasing the optical DC illumination
the noise decreases and approaches the shot noise limit.
The maximum SNR of an OCT system is reached when (1) the detection is shot noise
limited and (2) the detection bandwidth matches the interferometric signal bandwidth [20,21].
In the present experimental conditions we calculated the optimum SNR to be 92 dB.
However, the current implementation of the electronic filter integrated in each pixel (110x110
µm2) of the SPDA does not allow a very fine filtering of the signal. We estimated the filtering
bandwidth to about 40 kHz. The shot noise limited SNR of our setup is then 77.6 dB which is
close to the measured SNR.
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3.2. Sample images
In the following we reproduce image data that has been acquired during a 1600 ms time
interval at a rate of 25 volumes per second, i.e. a time sequence of 40 volumetric data sets
with dimensions of 58x58x58 voxels each. The raw data is in the form of a one-dimensional
array of 58x58x58x40 16-bit integers from which we have reconstructed the following two-
and three-dimensional images and movies.
Figure 2 shows a schematic of the imaged volume in relation to the sample and three
tomographic images along the sample symmetry axes at the beginning of the time sequence.
The image size format has been adjusted to represent the true geometric dimensions of
210x210 µm2 and 210x80 µm2, respectively. All images are shown in inverted grayscale
reflectivity coding. The en face image (center) shows the shadow caused by the hair at the
height of the supporting glass plate. On the longitudinal cut (right top), taken parallel to and at
the center of the hair strand, one distinguishes its upper and lower surfaces. The cross-
sectional cut (right bottom) exhibits the profile of the hair on the glass plate.
Fig. 2. (left) Schematic of the imaged volume (dashed parallelepiped) in relation to the sample
(hair strand on glass slide). (center) En face image (210x210 µm2) at the height of the contact
between hair and glass. (right top) Longitudinal cut (210x80 µm2) parallel to and at the center
of the hair. (right bottom) Cross-sectional cut (210x80 µm2) perpendicular to the axis of the
hair strand.
The reflectivity scale bar of Figure 2 also applies to all following tomographic images
which we represent for simplicity in a square 58x58 pixel image format. Figure 3 is a movie
that shows the temporal evolution of the sample in the three tomographic planes previously
discussed. During the first 960 ms (24 time frames) no modification of the hair is visible. The
incident laser radiation is highly absorbed and locally heats up the hair strand. At frame 25
(1000 ms after the beginning of irradiation) first modifications become noticeable in the
longitudinal and cross-sectional cuts. We see locally some higher reflectivity signals from
inside the hair volume, to the detriment of the signal corresponding to the lower side of the
hair that starts to be shadowed by these new backscattering sites. On the en face image these
changes are not yet visible, because the interaction between the laser and the hair has up to
this moment not progressed far enough in depth to modify the hair’s shadow on the glass
plate. On the next frame (40 ms later) a burn crater in the center of the hair can clearly be
distinguished on the longitudinal and cross-sectional cuts. Its depth can be estimated to
approximately 45% of the hair’s thickness. The formerly smooth surface of the hair becomes
very fragmented and scattering in the area of the burn crater and the signals from the lower
hair surface disappear because of shadowing. Reflectivity signals from above the former
position of the hair appear and might be identified as debris. The localization of the laser-hair
interaction to the center of the optical field is due to the Gaussian like spatial intensity profile
of the probing laser beam. Note that the SPDA cannot compensate for non-uniform
illumination since settings are common to all pixels. The area of interaction can be measured
to be of a diameter of about 120-130 µm. On the next frame (40 ms later) the hair’s
modifications become finally visible on the en face image. The hair starts to swell laterally
into a bulb and continues doing so for the next 520 ms until the end of the acquisition. This
bulb attains approximately 150% of the hair’s original width. The swelling can of course be
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observed on the cross-sectional images as well. However, other than the lateral swelling no
other major modifications are visible in this final phase of the interaction.
Fig. 3. (149 kB) Tomographic images acquired during a 1600ms time interval at a rate of 25
volumes per second (40 time frames). (left) En face image (210x210 µm2) at the height of the
contact between hair and glass. (center) Cross-sectional cut (210x80 µm2). (right) Longitudinal
cut (210x80 µm2) parallel to and at the center of the hair. Reflectivity grayscale as in Figure 2.
Because of an experimental problem of synchronization images of even and odd volumes
are slightly shifted in relation to each other, which causes the image jumps on this movie. This
imperfection can easily be corrected for and this has been done in Figure 4. Furthermore, the
last column of photodetectors (rightmost column of pixels on an en face image) exhibits a
markedly different response than the others, which causes image artifacts. This is due to a
different electronic pixel layout that has been realized on the last column for experimental
purposes. The distortions visible on adjacent columns in the beginning of the time sequence
(clearly visible as trailing signals of the glass plate on the cross-sectional cuts) might be
related.
Figure 3 illustrates three tomographic views that permit to visualize relatively well the
dynamic phenomenon observed. However, much more data has been acquired with this 3D
OCT method and any chosen view could be visualized. In order to allow an inspection of the
entire volume at one glance we use the whole data set to generate a three-dimensionally
rendered representation based on isosurfaces. Each of the 40 time frames has been rendered as
described in [22] and combined into the movie shown in Figure 4. This representation goes far
beyond tomographic images and is very useful for localizing regions of particular interest.
Fig.4. (740 kB) Movie of a 3D rendering of the sample based on isosurfaces. To facilitate the
comprehension of this particular perspective we indicate the situation of the hair and the glass
slide by the colored lines in the first frame.
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4. Conclusion
In conclusion, we have implemented a parallel OCT system capable of 3D data acquisition at
video-rate. The key element of the system is a smart pixel detector array, conceived and
developed specifically for en face OCT imaging. Combined with a femtosecond light source
and a microscopic imaging scheme this system allows for both high longitudinal and
transverse resolutions when limited to a small sample volume. We have illustrated its
performance by imaging the time-resolved thermal damage of a strand of dark human hair
under the influence of the probing laser beam. Besides tomographic images we have also
shown a three-dimensionally rendered movie.
Acknowledgments
We would like to thank P. Thevenaz for his help in the 3D rendering of our data and S.
Bourquin for valuable discussions.
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