Abstract
Air Force
Portable Device for Retinal Imaging
Military aircrew members, or anyone involved in a dangerous activity requiring very high
visual performance, would benefit from a field-deployable screening device to detect
lesions of the retina before they could cause permanent loss of visual sharpness or
blindness. At present, there is no accepted diagnostic or screening protocol for ocular
lesions and retinal cell loss associated with retinal laser injuries. Photographic fundus
imaging is a standard procedure to evaluate ocular lesions; however, it can only provide a
two-dimensional retinal surface image. The newly developed OCT technology can provide
three-dimensional retinal images that will greatly facilitate the in-field diagnosis and
treatment of ocular lesions caused by laser weapons or laser guided aiming devices. This
technology uses a swept laser source for Fourier domain optical coherence tomography
(FDOCT), with 1 m in tissue band probe light, high-contrast and three-dimensional for
retinal imaging. This SS-OCT has a measurement speed of 20,000 A-lines/s, a depth axial
resolution of 12 m in tissue and sensitivity of 99.3 db, and coherence length of 9.8 mm at
2 kHz sweep rate. A software-based algorithm of scattering optical coherence angiography
(S-OCT) was used to visualize the 3-D retina of the human eye. Based on the swept source,
an FDOCT system was developed that can achieve 12 m axial resolution in tissue.
Imaging of a human retina was demonstrated with the FDOCT system, and the FDOCT
resolution was improved by reducing the effects of the tails or side lobes using the wiener
filter logarithm that was implemented in the S-OCT.
Abdelhamid Jnane
City College Of New York
CUNY
Air Force
Portable Device for Retinal Imaging
Mentors: Dr.zhongping Chen
Dr. Qiang wang
University of California, Irvine
Beckman Laser Clinic
Key words
Optical Coherence Tomography (OCT)
Fourier domain OCT
Spectral domain OCT
Swept Domain OCT
Scattering OCT
Interferometer
Full-width at half maximum (FWHM)
Gaussian Spectrum
Wiener filter
Introduction
Optical coherence tomography is a non-invasive in vivo imaging technique, which has undergone
rapid development since its invention in 1991 [1,2]. In 1991, optical coherence tomography
(OCT) was reported for the first time [3]. Using a Michelson interferometer and a low coherence
light source, OCT performs cross-sectional imaging of internal tissue microstructure by
measuring the echo time delay of backscattered infrared light [4,5]. In an interferometer, light
emitted from a source is divided by a beam splitter in two directions towards two mirrors, a
reference mirror and a sample mirror. The back-reflected beams recombine at the beam splitter
and are guided to a detector. When the reference arm is moved while the sample is fixed, an
interference pattern is measured by the detector. Interference, however, will only be detected
when the difference in optical distance traveled by the light in both arms is less than or equal to
the coherence length of the light source [6,7]. The interferometer output is detected and
demodulated to measure the echo time delay and magnitude of the backscattered light. By
acquiring sequential transverse axial measurements and removing motion artifacts using crosscorrelation algorithms [8], the resulting data set is a two-dimensional array representing the
optical backscattering in a cross-sectional plane of the tissue. This can be digitally processed and
displayed as a gray-scale or false-color image [9].
There are two types of OCT. The first kind is called time domain OCT (TDOCT) [10], which is a
single longitudinal scattering profile that translates the reference mirror. In this system, both the
depth and transversal directions have to be mechanically scanned to register a two-dimensional
cross-section map of the region of interest within the object [11,12,13]. Fourier domain OCT
(FDOCT) is an alternative to TDOCT, which provides high speed and high sensitivity [14,15]. In
FDOCT mechanical depth scans are not required. The reference mirror position is fixed and the
spectrum of the light at the output of the Michelson interferometer is recorded by a spectrometer
equipped with a high-speed multi-element detector. To provide the profile of the sample
reflectance versus depth (optical A-scan), the Fourier transformation of the interference pattern is
used. FDOCT detection can be performed in two ways: spectral domain OCT (spectral OCT),
using a spectrometer with a multi channel analyzer [16,17], or swept source OCT, using a rapidly
tunable laser source [18,19,20]. Spectral OCT and swept source OCT have been used for
ophthalmic imaging in the anterior eye and retina and have demonstrated ultrahigh-resolution
imaging and high speeds when compared to TDOCT [21].
The longitudinal resolution in OCT is proportional to the source spectral bandwidth, which
governs the axial point-spread function. Therefore, the full-width at half maximum (FWHM) of
the interference fringe envelope is used to define the axial resolution of an OCT system. Side
lobes and tails in the interference fringe envelope are usually caused by the non-Gaussian spectral
shape, due to several types of sources (e.g., semiconductor optical amplifier) of light. OCT
resolution can be improved by reducing the effects of the tails or side lobes using some signalprocessing algorithm. The Wiener Filter logarithm was implemented in this study. [22].
FDOCT with a wavelength of 1.3 m is widely used in many ophthalmology applications because
it provides higher penetration for a tissue than the one with an 830 nm wavelength. [23].
However, the 1.3 wavelength is not suitable for the investigation of retinal imaging because the
water absorption of 1.3 m light is 38 times greater than that of 830-nm light [24].
Because water absorption has a local minimum at 1.06 m, wavelength with 1 m is the
alternative solution for high penetration retinal OCT. The 1 m wavelength can easily access and
travel in the vitreous body of the retina. Because SD-OCT requires a high-speed spectrometer, it
is not suitable to assemble an SD-OCT with a 1 m probe. An SS-OCT scheme is the alternative
for 1 m FDOCT because of its higher speed and higher tolerability in sampling the signal [25].
In this study, we describe the development of new FDOCT device combined with a 1 m swept
source light to visualize 3-D retinal structure imaging. We also demonstrate a novel
implementation of numerical spectral shaping using Wiener Filter algorithm to smooth the
spectrum, and we conduct a preliminary testing of our system.
Fig. 1. Schematic of the FDOCT system: Collimator; adjustable neutral density attenuator;
D1 and D2; photodetectors.
Spectrum envelope of the source
80
Axial PSf without spectral reshaping
200
Amplitude[arbitrary units]
Intensity [Arbitrary units]
70
60
50
40
30
0
20
1020 1030 1040 1050 1060 1070 1080 1090 1100 1110
Wavelength [nm]
0
100
200
300
400
Miror displacement[mic.m]
Fig. 2. (a). Non-gaussian Spectrum from the swept light source registered by the OCT
spectrometer. (b). The curve represents the spectral uncorrected response for Coherence
envelopes determined from glass slide Reflection.
Method and System Hardware:
The scheme of our FDOCT instrument (Fig.1) uses a 1 m swept laser source as a light source
with booster and without booster (2 kHz repetition rate; Peak power of 0.7 to 4 mW; FWHM_
sweep bandwidth of 31 nm centered at 1078 nm without booster and 1061 nm with booster). The
hardware specifications used in this OCT design were as follows:
Collimator (OFR Inc., 1060 nm); Gold mirror (Thorlabs Inc., ½”); 2 circulators (Agiltron Inc.,
1060 nm); 70/30 2x2 couplers (AC Photonics Inc. 1060 nm; 50/50 2x2 couplers (AC Photonics
Inc. 1060 nm); X stage (Newport Inc.); Light source (Santec Inc., 1060 nm, 28 kHz); Aluminum
case (Hammond Manufacturing, 10.03” x 9.63” x 3.84”). The two 50/50 fiber couplers were used
to increase SNR of the system. The output light from the swept source was directed into reference
and sample arms by a 2*2 coupler. An adjustable attenuator was used to attenuate the reference
power for enhanced sensitivity in the reference arm. The imaging speed of the system is 20 k Alines/second and the power delivered to sample arm is 1.2 mw.
The signal from the balanced detector was converted by a data acquisition board sampling at 10
MHz. The number of data points for each A-line data acquisition during the frequency scan was
1024. The axial resolution in free space was determined to be 17 m corresponding to the
effective axial resolution of 12 m in tissue by measuring the point spread function with a partial
reflector at a depth of 900 m.
Numerical spectral reshaping:
As shown in Fig. 2 light source has a non-gaussian shape caused by optical fiber components of
the light source. To get a gaussian shape of the spectrum, we applied a numerical method by using
a Wiener Filter. To design a Wiener Filter, it is essential to determine the envelope of the
wavelength-dependent fringe contrast. First, we extracted the wavelength-dependent fringe
contrast from the OCT spectral signals.
The spectral interference signal can be expressed as follows:
S j (k) = Se (k) cos j (k)
(1)
Where Se (k) is the contrast, cos j (k) is the fri
j (k) is the phase of the j-th fringe.
The ensemble average of S j(k)2 in j can be rewritten as follows:
S j (k) 2j = Se (k)^ 2 cos^2 j (k)j
(2)
Since the squared cosine term can be reduced to a constant value of ½, the contrast envelope is
rewritten as:
–Se (k) =√ (2/N*Σ (j) Sj (k)^2
(3)
Substitution of equation (3) in a Wiener Filter yields the desired spectral reshaping filter that
will be implemented in our OCT software, as illustrated below:
W (k) =(Se (k)/(Se (k) ^2 +Nc))*Gaussian (k)
Where Nc is a constant depending on the SNR of the detection system and Gaussian (k) is a
Gaussian window to reshape the spectrum to a Gaussian profile.
For the Wiener Filter implementation for our OCT, we wrote a C++ program portion to reshape
our spectral light source and obtain a smooth gaussian shape of the signal. For the spectral
envelope calculation we used the following polynomial 8 degrees:
Spectral Envelope = A + B1*j + B2*j^2 + B3*j^3 + B4*j^4 + B5*j^5 + B6*j^6 + B7*j^7+B8*j^8
Where A=23.84411; B1=0.95364; B2= -0.0062; B3=9.5098E 5; B4= -7.19185E 7;
B5=2.49286E-9; B6= -4.32256E-12; B7=3.66599E-15;B8=1.21285E-18.
We then express the Gaussian envelope and spectral-reshape as:
Gaussian Envelope = 100*exp (-(j-425.0)* (j-425.0)/240.0/240.0)
Spectral Reshape=Gaussian Envelope [(2.0*Spectral-Envelope +constant factor)];
Results and discussion
Fig.3. A compact FDOCT portable device for retinal imaging
1.002
70
Specrtrum envelope of the source
Axial PSF with spectral reshaping
60
Amplitude[arbitrary units]
Intensity [arbitrary units]
1.000
0.998
0.996
0.994
0.992
50
40
30
20
10
0
0.990
1020 1030 1040 1050 1060 1070 1080 1090 1100 1110
Wavelength [nm]
0
100
200
300
400
Miror displacemen[mic.m]
Fig. 4. (a) Gaussian Spectrum from the swept light source registered by the OCT
spectrometer. (b) The curve represents the spectral corrected response for Coherence
envelopes determined from glass slide Reflection.
The rat animal model will be used for in vivo imaging of retinal damage to evaluate the proposed
field deployable imaging device. A mode locked Titanium: Sapphire laser with 800 nm
wavelength, 76 MHz repetition rate and 170-femtosecond pulse or Continuum MiniliteTM,
Q-switched Nd: YAG laser of 532 nm wavelength, 1~15 Hz repetition rate and 3~5 nanosecond
pulses will be used to create laser lesions in rat retina. We will use the ophthalmoscope threshold
of 5.9mJ (Robert J. Thomas, Lasers in Surgery and Medicine 31:9-17(2002)) as standard lesion
dosage. Three types of lesions will be created by exposing to 100%, 200% and 300% of the
standard lesion dosage. For the exposure time of 0.25 s (controlled by an electric shutter) that will
be used in our experiment, the laser power for the three different dosages will be 23.6 mw, 47.2
mw and 70.8 mw respectively. Three groups of animals (Group A created with 23.6 mw, 0.25 s;
Group B created with 47.2 mw, 0.25 s; Group C created with 70.8 mw, 0.25 s) will be required. A
total of five animals will be assigned to each group to evaluate the time response of the retinal
lesion at each of the five time points: 0 h, 1 h, 5 h, 24 h and 48 h post injury. A total of 15 rats
(three groups x five time points/group x one rat/time point) will be needed for this pilot study. To
accommodate factors such as laser misalignment during the creation of retinal lesions, the same
laser parameter will be used to create wounds on both eyes of the rat. The rats will be housed
individually and the cage for each rat will be tagged to identify its group and lesion response time.
The following procedures describe how the experiment will be performed on each rat.
After general anesthesia, retinal lesions on both eyes will be created and then imaged at
the time point assigned to the animal. When the post injury time point for imaging is
greater than 1 hr, the animal will be recovered from anesthesia and then reanesthetized at
the designated imaging time point. Although the retinal injury will cause no pain, the
animal will be monitored for signs of squinting, scratching and rubbing of eye as well as
the swelling of the eyes during the time intervals between retinal injury and imaging time
point. For those animals with long response times (24 hr or 48 hr), the animals will be
transported back to the animal housing facility. At the assigned response time, the animal
will be transported back to the BLI-S306, reanethetized, and the photographic fundus
images and the 3-D retinal optical coherence tomography images will be obtained for both
eyes. Immediately after the imaging procedure, the animal will be euthanized without
recovering from general anesthesia and the eyes will be harvested. For each anima,l the
experimental end point is reached at this time.
In the pilot experiments described above, a total of 30 photographic fundus images and 30 3-D
OCT images will be documented for the 15 rats. Two groups of ophthalmologists will evaluate
the images. The first group will evaluate only the photographic fundus image to make an
evaluation of the retina injury. Their evaluations will be matched, compared and scored according
to the histology evaluation of the specimen. The second group will evaluate both the photographic
fundus images and 3-D OCT images to evaluate the retina injury. The evaluation will be matched,
compared and scored to the histology evaluation. The improved effectiveness of the retinalscreening device that uses both fundus photography and OCT imaging can be derived as the ratio
of the total scores of the second review group to the scores of the first review group.
Conclusion
We have developed a FDOCT device with a 1 m swept light source, which has the following
specifications: Dimensions of 8” x 8” x 4”; sensitivity of 99.3 db; High speed (20,000 A-lines/s);
High resolution (12 mm); coherence length of 9.8 mm at 2 kHz sweep rate.
We have also shown that FDOCT resolution can be improved by reducing the effects of the tails
or side lobes by implementing the Wiener Filter algorithm to get a gaussian shape of the spectral
light source. The algorithms we have used take advantage of the convolution property of Fourier
transformation. Therefore, extensive computation that can slow the speed of the OCT is not
necessary
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Acknowledgements
Dr. Zhongping Chen
Dr. Qiang Wang
University of California, Irvine
Beckman Laser Institute
IM-SURE Program
Dr. Said Shokair
NSF