Whole Eye Optical Coherence Tomography (OCT) to Improve Refractive Surgery and Eye Care
Conditions: Image, Body
Sponsor: Duke University

 
 
 
 
 
 
 
Whole Eye Optical Coherence Tomography (OCT) to Improve Refractive Surgery and Eye Care 
Document Date: August 14, 2014 
NCT03219567 

 
Figure 1. Proof of principle 
demonstration of 
differences in peripheral 
retinal thickness 
measurements using OCT. 
The OCT beam is shown in 
red pivoting through the iris. 
The eye model is based on 
published schematic eyes 
[52,58]. The desired 
thickness is the surface 
normal thickness h. The OCT 
measured thickness is along 
the A scan in red. The 
difference between h and the 
OCT measured thickness 
increases as the half scan 
angle β increases. 
SIGNIFICANCE 
Optical coherence tomography (OCT) is a non-contact, micrometer scale imaging technique which is 
widely used in clinical eye care and research. In 2010, it was estimated that over 16 million clinical ophthalmic 
OCT procedures were performed in the United States alone [5]. This averaged one OCT performed on a 
patient’s eye every 2 seconds in the U.S., not even including research OCT usage. 
OCT is widely used because in contrast to other clinical imaging techniques, it provides clinicians and 
researchers with high resolution in vivo images sufficient to visualize layered microanatomy. Initially, these 
images were used qualitatively to describe features and note presence or absence of pathology. Simple 
quantitative measures such as layer thicknesses of the cornea or the retina were introduced subsequently [6]. 
More sophisticated quantitative geometric measures such as curvature, slopes, and other volumetric 
descriptors of shape in the cornea, retina, and optic nerve head have been recently described [7-9]. OCT 
images, however, are subject to geometric, refractive, and motion artifacts during acquisition [1-4]. This results 
in spatial distortions in OCT images of the eye, and affects even simple measurements on that representation 
of the eye (Fig. 1). Though current OCT systems have small fields of view (10°) that limit the effect of some of 
these distortions for simple measures, quantitative geometric measures will be increasingly impacted by the 
rapid development and imminent advent of wide-field (40° to nearly 90°) OCT systems both in research and 
commercially [10-12]. 
To faithfully 
represent the eye 
using OCT requires 
a full understanding 
of 
the 
distortions 
introduced by the 
entire optical train 
from 
the 
OCT 
system through the 
eye’s own optics, 
particularly 
for 
retinal imaging [1-
3,13-16]. 
Because 
of light source and 
optical 
limitations, 
current clinical OCT 
implementations 
cannot image the 
entire 
eye 
in 
a 
single acquisition (despite a Photoshop-ed “whole eye” OCT in a marketing brochure [12]). A few research 
groups have begun to address this limitation using multiple reference arms [17] or vertical cavity surface 
emitting laser (VCSEL) sources with meter-scale fall-off [18] to achieve axial imaging ranges sufficient to image 
to the back of the eye. Our group was the first to publish the combination of a technology with sufficient axial 
range for whole eye OCT (coherence revival) combined with dedicated polarization-encoded dual-channel 
optics designed for optimal imaging of each depth region [19]. Without these dedicated optics, the lateral 
imaging range and resolution of ocular images from OCT are compromised, regardless of the axial range of 
the system [18]. Additionally, our group has been among the leaders in developing algorithms to remove 
refractive and geometric distortions in both anterior and posterior eye images to improve the quantitative 
accuracy of OCT representations of the eye, particularly for diagnostic measurement of morphology [1,4,20]. 
In this proposal, we will develop the hardware systems and software algorithms necessary to enable 
simultaneous OCT imaging of all the refractive surfaces of the eye. We will apply these technical advances to 
two specific clinical applications where accurately measuring the surface geometry is an important diagnostic 
task. These clinical aims will also validate the OCT measurement accuracy throughout the eye. 
In the front of the eye, enabling accurate measurement of the corneal shape after LASIK and other 
laser refractive surgeries in over 12 million patients who have had these procedures is a recognized research 
priority of the National Eye Institute [21]. The inability of current diagnostics to accurately measure the post-
surgical corneal shape has resulted in unexpected hyperopia after subsequent cataract surgery [22]. For these 
patients who were previously accustomed to good spectacle free vision, post-surgical hyperopia means 
complete dependence on glasses for all distances. A large number of compensatory algorithms have been 

FIGURE 2. Ultra-wide field swept-source retinal OCT at 1060nm. Results 
from prototype system imaging with 85° FOV on the retina relative to the 
posterior nodal point of the eye. Representative B-scan (A) and summed voxel 
projection (B) are shown. Layered retinal structures are visualized over the 
complete field of view. 
offered for this measurement problem, and the net result clinically is that cataract surgery planning for this 
population is confusing and outcomes are unpredictable [23]. This unpredictability is difficult in a clinical 
environment where patients routinely expect 20/20 spectacle free vision after refractive and cataract surgeries. 
In the back of the eye, measuring surface geometry is also diagnostically important (please see 
supporting letter from Collaborator Terri Young, M.D.). The standard simple axial length measure has proven 
inadequate for understanding the myopic (near-sighted) eye and its development. Specific three-dimensional 
alterations in the shape of the posterior eye in those with myopia have been correlated with progression of this 
disease [24] and with vision threatening pathology [25,26]; hence, posterior eye shape could be used for risk-
stratification with appropriate follow-up and counselling. Recently, off-axis peripheral treatment of refractive 
error (not standard axial refractive correction) has demonstrated promising arrest of myopia development [27]. 
Though it would be reasonable to measure posterior eye shape with long available ophthalmic ultrasound, a 
search of the literature (PubMed using “eye” AND “shape” AND “myopia” AND “ultrasound”) returned only one 
paper using ultrasound for this purpose [28]; the remainder used ultrasound only to measure axial length. 
Instead, posterior eye shape is currently obtained using MRI [24-26,29-31] with its attendant cost and logistical 
barriers. Lowering the cost and logistical barriers to obtaining the shape of the posterior eye using ubiquitous, 
non-contact OCT would allow more widespread clinical use of this measurement for risk stratification of 
myopes and for study of myopia progression. Further, using OCT provides a 10-fold increase in resolution over 
posterior segment ophthalmic ultrasound for measurement accuracy and also provides the layered 
microanatomy that is used widely by clinicians and researchers. 
Our contribution will be significant because we will address the important problem of distortions 
inherent to widely used ophthalmic OCT that impact quantitative measurements from those images, particularly 
with imminent wide-field implementations. 
Our proposed project will improve clinical care by two immediate clinical applications. First, measuring 
post LASIK corneal power and anterior segment biometry will improve cataract surgery outcomes in this large 
population, thus reducing potential physical and financial costs associated with unpredictable visual outcomes. 
Second, by using OCT to recover the shape of the posterior eye, our proposed project makes this diagnostic 
biomarker widely available clinically and for research. OCT is already used clinically in patients ranging from 
infancy (without sedation) to adulthood, and would be significantly less costly than MRI for this purpose. 
Achieving our aims will change ophthalmic diagnostic imaging by enabling widely used OCT to produce 
accurate, faithful representations of the imaged eye as a whole with broader implications for improving the 
accuracy and reliability of any quantitative measurements based on OCT images.  
 
INNOVATION 
 The proposed research is innovative 
because it challenges the conventional 
representation of the eye in OCT images 
by recognizing the eye as a complete 
optical system, and not as non-interactive 
anterior 
and 
posterior 
regions. 
By 
considering 
the 
image 
forming 
contributions of the whole eye, our OCT 
hardware and software developments will 
produce 
quantitatively 
accurate 
representations of both the anterior and 
posterior eye. The expected innovations 
include: 
1. Novel optomechanical hardware designs to enable simultaneous imaging of both the front and back of the 
eye with focused fields of view larger than current systems dedicated to only the front or the back of the 
eye (Fig. 2). 
2. Hardware and algorithm/software developments to remove the effects of motion and optical artifacts that 
corrupt the accuracy of anterior eye OCT. This will enable three-dimensional anterior segment biometry 
with OCT for clinical use in surgically and non-surgically affected corneas. 
3. In addition to the above, algorithmic developments to remove the effects of optical artifacts that distort the 
spatial accuracy of posterior eye OCT. This will enable three-dimensional posterior segment biometry with 
OCT for clinical and research use in myopia and other posterior eye applications. 

FIGURE 3. Simultaneous OCT imaging of the anterior 
and posterior eye from our previous work [19]. Both 
volumetric renderings (left) and representative constitutive 
B-scans (right) are shown. Of note, the FOV and 
resolution of the simultaneously acquired images are both 
comparable to images acquired from dedicated anterior or 
posterior eye OCT systems. (The anterior and posterior 
eye images do not share the same scale.) 
Together, these innovations will produce a spatially accurate portrayal of the eye from ophthalmic OCT and 
provide a solid foundation for all quantitative measurements on the imaged eye. Our previous studies which 
have laid the groundwork for all of these innovations [19,32,4] strongly suggest that these advances will 
improve the diagnostic accuracy of OCT derived measurements for both the front and back of the eye. This 
departure from the status quo is expected to enable the widely available OCT platform to faithfully represent 
the imaged eye. 
 
APPROACH 
Specific Aim 1: Develop hardware modifications to OCT to allow simultaneous imaging of both the 
anterior and posterior eye. 
SA1. Introduction. To produce images of the eye 
accurate enough to perform surface metrology, we require 
images containing all of the important optically refracting 
surfaces in both the anterior (cornea, lens) and posterior 
eye (retina). By directly imaging both the anterior and 
posterior eye, optical analyses of the eye can then be 
performed directly on the imaged optical surfaces. To 
produce a sharply focused OCT image at both the anterior 
and posterior eye requires deliberate design because the 
imaging paths to these two locations are optically different. 
The objective of this aim is to develop the 
optomechanical modifications to the usual OCT topology 
that will enable, at minimum, focused standard field of view 
(FOV) images of both the anterior and posterior eye 
simultaneously. 
Simultaneous 
capture 
ensures 
co-
registration of the anterior and posterior images to produce 
single, unified representations of the eye. 
The rationale for this aim is that successful 
completion of this aim will produce OCT images of the 
major ocular surfaces that are optically accessible. Our expectation is that the image volumes produced by this 
system will contain the information necessary to improve the accuracy of the represented eye for subsequent 
optical analysis, image distortion correction, and diagnostic use. 
SA1. Justification & Feasibility. While there have been prior implementations of simultaneous whole eye 
OCT systems [14,15], those implementations were intended primarily to measure distances to various 
surfaces. These systems utilized telecentric scanning focused on the anterior eye. Posterior eye images from 
these systems had extremely limited FOV due to telecentric scanning at the anterior eye; in fact, when 
refraction corrected, the retinal FOV from these systems was an oversampled small spot. These systems also 
lacked lateral resolution at the macula because the system focus was designed for the anterior eye. While 
these limitations do not hamper measurement of distances, dedicated wider fields of view focused on the retina 
are required to truly study the posterior eye. 
Our team has developed the component technologies for an OCT system that is capable of 
simultaneous, focused imaging of both the anterior and posterior eye [19]. In this system, the OCT light is 
divided using polarization into anterior and posterior eye channels. Both paths independently image their 
respective area of the eye simultaneously. On return, the coherence revival technique unique to specific swept 
source lasers is used to encode the different imaged depths [33]. This provides both correction of the complex 
conjugate artifact for extended imaging range in the anterior eye, as well as depth multiplexing with separation 
of the anterior and posterior eye scans within the final images. An image volume of a living subject’s eye 
obtained with this system after post-processing is seen in Figure 3. 
While this system demonstrated the feasibility of simultaneous dual depth imaging with OCT using only 
a single source and detector, this demonstration system will require additional design to achieve the goals of 
this proposal. Most notably, the anterior eye channel of this system was not designed to image the anterior eye 
from the cornea to the lens, and the posterior eye FOV was less than the usual 30° in standard ophthalmic 
photography. With this demonstration system as a basis, however, we will make informed design modifications 
to improve this system for the stated aims of this proposal. 
SA1. Research Design. Whole eye OCT system design. The basic OCT configuration consists of the source, 
detector, sample arm, and reference arm. To meet the clinical needs of this project, we will substantially 

FIGURE 4. General schematic of sample arm design for simultaneous OCT imaging of the 
anterior and posterior eye (similar to our previous design in [19]). This design relies on 
polarization encoding to separate the light source (purple) into anterior (blue) and posterior (red) 
eye paths. Dedicated optics for the anterior and posterior eye paths allow for optimal imaging of 
each location even through acquisition is occurring simultaneously.  In the first specific aim, we will 
redesign: (1) anterior eye optics for increased depth of focus, (2) posterior eye optics for increased 
FOV, and (3) galvanometer pair to include a 4F, 1x relay. L1: collimating lens, HWP: half wave 
plate, FM: fold mirror, PBS: polarizing beam splitter, GY/GX: 1D galvanometers, L2: 4F telescope 
with1X magnification to image GY onto GX, L3a-b: retinal telescope (L3b is adjustable to optimize 
focus at the posterior eye), L4a-b: anterior eye optics, DCM: dichroic mirror for iris camera. 
redesign the sample arm (Fig. 4). There will be three major changes: 1. Redesign of the anterior eye channel 
optics to enable imaging from the anterior cornea to the posterior lens, 2. Redesign of the posterior eye 
channel optics to increase the FOV, and 3. Redesign of the galvanometer pair configuration to ensure uniform, 
symmetric scanning. An iris camera will also be added to aid clinical use and alignment. 
The anterior eye 
channel 
requires 
redesign 
to 
enable 
imaging from the anterior 
cornea to the posterior 
lens. Nominal values for 
the 
component 
thicknesses in a 60 year 
old are as follows: cornea 
thickness 
0.5 
mm, 
anterior chamber depth 
2.86 
mm, 
and 
lens 
thickness 4.24 mm [34]. 
Together this suggests 
that an imaging depth on 
the order of 7.6 mm is 
required for full anterior 
eye imaging. For OCT, 
there are both signal 
(falloff) and optical considerations that affect imaging depth range. From a signal standpoint, the 6 dB falloff 
range for our prior coherence revival SS-OCT system was 8.5 mm [19]; this exceeds our nominal requirement. 
We also expect a slight increase in this range with increased digitizer bandwidth in the newer planned system. 
Optically, we need to increase our depth of focus (DOF) to image through the desired depth range. 
Increasing DOF, however, will decrease the lateral resolution. To determine the appropriate balance between 
DOF and lateral resolution, we consider our requirements. Commercial anterior eye OCT systems with sources 
1000 nm or greater have reported resolutions of 10-18 µm axially and 30-60 µm laterally. Using the minimum 
of these parameters as guidance, we have created an initial anterior eye sample arm design with a DOF of 6.7 
mm and a theoretical full width-half max lateral spot size of 39.5 µm in air. Coupled to our 1060 nm SS-OCT 
engine with a measured axial resolution of 7.8 µm in air, we were able to produce an image of the entire 
anterior eye with this preliminary sample arm design (Fig. 5). 
Regarding the posterior eye channel, redesign is required to increase the FOV to measure posterior 
eye shape. Figure 2 shows the imaging results from a preliminary posterior eye sample arm built with large 
diameter double aspheric optics which produces 85° FOV. This is nearly triple that of standard ophthalmic 
photography, and all major posterior ocular features (optic nerve, macula, vascular arcades) are visible while 
still retaining high axial resolution as evidenced by presence of layered microanatomy in the B-scan. Also the 
system is implemented at 1060nm, enabling penetration through the full choroidal depth. 
 
Together, the re-designed anterior and posterior eye imaging sub-systems meeting our design 
requirements shown in Figs. 2 and 5 will be incorporated into the polarization-encoded, coherence-revival 
based system depicted in Fig. 4 to accomplish true whole eye OCT – simultaneous anterior and posterior eye 
OCT with optimal focusing and large fields of view at both locations – for the first time. 
 
The light source for this whole eye OCT will be the same commercial external-cavity tunable swept 
source laser operating with a nominal center wavelength of 1060 nm, 100 nm tuning bandwidth, and sweep 
rate of 100 kHz (Axsun Technologies) used in our previous work. A balanced receiver detection design will be 
used with a 1 GHz InGaAs balanced receiver (WL-BPD1GA, Wieserlabs) and digitized at 1.8 GS/s on a 12 bit 
(9.1-ENOB) digitizer (ATS9560, Alazar Tech). This digitizer has larger bandwidth (800 MHz) than the original 
system (500 MHz) which will decrease fall off and increase SNR farther from DC. Otherwise, the detection 
design reflects what was used in our earlier system to achieve dual depth, coherence revival imaging. 
The complete system will be assembled into a rack mountable breadboard and case to allow use in the 
clinical environment without exposed optical elements. The sample arm will be mounted onto an adjustable, 
portable table with head rest and chin rest for patient comfort and to minimize gross patient movement. The 
reference mirror will be on a ruled translation stage to allow calibration of system to eye distance as described 

FIGURE 5. (A) Increased depth range anterior eye sample arm and (B) resultant 
image quality in a human subject. This sample arm design has fields of view of 17 x 
17 mm, depth of focus of 6.7 mm, and a theoretical lateral full width-half max spot size of 
39.5 µm. The resultant image shows both the anterior and posterior cornea and the 
anterior and posterior lens. Vertical artifacts from increased light reflection at the iris and 
sclera are present in this iteration of the system design. Within the pupillary zone 
(“optical zone” of the eye), these brighter artifacts are not present. The OCT engine is 
the same as used in our prior work [19]. 
below. 
The 
system 
will 
be 
controlled by custom software 
written in C++ utilizing GPU 
coding for real time OCT data 
processing and display at 100 
kHz A-scan rate. 
System 
calibration 
for 
quantitative analysis. Once the 
system 
is 
assembled, 
the 
sample space of this system will 
need to be characterized for 
subsequent 
quantitative 
analyses as we have done in 
past work [20]. There are three 
primary quantities which will need to be calibrated: distance of system to the sample (eye), lateral pixel pitch, 
and axial pixel pitch. The lateral and axial pixel pitches define the sample space dimensions within the digital 
image and is determined by imaging a dual-axis linear-scale stage micrometer (NT58-763; Edmund Optics). 
The system to eye distance affects magnification and telecentricity of the scan and can be determined 
from the zero path length delay (ZPD) in the anterior eye image. The location of the ZPD in the image is 
equivalent to the reference mirror position in Fourier domain OCT. To calibrate the system, we will image a flat 
target in air at a known, measured distance (close to the working distance) from the final optic of the sample 
arm. We will then adjust the reference mirror until the image of that flat target is at the ZPD within the anterior 
eye channel image. At this calibrated reference position, ZPD in the image represents a known distance from 
the final sample arm optic within the image. During clinical imaging, the distance from the final sample arm 
optic to the cornea (eye) will be measured by adding the known distance to the ZPD (calibrated) to the 
distance between ZPD and the corneal apex in the image. Any reference arm translation from the initially 
calibrated position is simply measured from the translation stage and added to the ZPD to cornea distance to 
determine the final system to eye distance. In summary, ZPD in image (ZPDimage) = known calibrated distance 
from final optic of system, ZPD to cornea (cornea-ZPD) = measured from image, reference arm translation 
from initial calibrated position (Δref) = measured from translation stage. The final system to eye distance = 
ZPDimage + (cornea-ZPD) + Δref. Distances to the posterior eye image content will be determined by measuring 
the fixed one laser cavity length offset from the anterior eye image used for coherence revival [19,32]. With 
system to eye distances known, we will accurately account for magnification and non-telecentricity optical 
artifacts affecting the imaged ocular structures. 
SA1. Expected Outcomes. This aim will produce an OCT system capable of simultaneous clinical imaging of 
the anterior and posterior eye. For the anterior eye channel, the resolution will be sufficient to measure corneal 
power and the imaging range sufficient to image from the front of the cornea to the posterior lens. For the 
posterior eye channel, the FOV will nearly triple standard ophthalmic photography (85° FOV relative to nodal 
point) with resolutions comparable to current retina-only OCT systems. As a whole eye system, both channels 
will be imaging simultaneously to produce coregistered volumes of all meaningful surfaces of the eye. The 
sample space of this system will also be characterized sufficiently to perform clinical surface metrology on the 
images of the eye. With this information, accurate clinical measurements for anterior segment biometry and 
posterior eye shape recovery can be performed. 
SA1. Potential Problems & Alternative Strategies. While we are confident that the coherence revival 
imaging strategy proposed above is readily extendable for whole eye imaging, it is possible that by the time 
this proposal is funded, alternative technologies to coherence revival such as a 1050 nm VCSEL source may 
become commercially available which would simplify the required technology development. If so, we will take 
full advantage of this by incorporating this source with our dual depth sample arm optics to obtain full fields of 
view for the posterior segment; otherwise, the VCSEL source alone only scans a small spot posteriorly [18]. 
Finally, we also recognize that there are ambiguities inherent to image reconstruction in the presence of 
the phakic (natural) lens. This is specifically addressed in the posterior eye reconstruction aim (SA3.Research 
Design and SA3.Potential Problems). In short, we will take a systematic approach starting with imaging of 
pseudophakic subjects (artificial lens of known optical prescription) for validation of the reconstruction 
algorithms followed by imaging of phakic subjects with reconstructions utilizing published gradient index 
models of the natural lens. 

 
FIGURE 6. Comparison of corneal refractive power 
measurements between motion corrected distributed 
scanning OCT (DSOCT), top, and the clinical standards 
(topography, bottom left and Scheimpflug photography, 
bottom right) [32]. In these normal corneas, the mean 
difference between OCT and the two clinical standards is 
less than 0.10 D with a standard deviation approaching the 
clinical refraction resolution of 0.25 D.  
Specific Aim 2: Develop and validate algorithms to produce accurate three-dimensional anterior eye 
OCT reconstructions for use in post LASIK cataract surgery patients. 
SA2. Introduction. Clinically measuring the refractive power of the post-LASIK cornea has proven to be a 
challenge with current diagnostics; this has led to unpredictable cataract surgery outcomes for these 
individuals who previously enjoyed high quality spectacle free vision after LASIK. The post-LASIK cornea 
violates assumptions made by current diagnostics to determine whole corneal power from a single 
measurement of the anterior corneal surface. While corneal power is thought to be the predominant factor in 
this problem, modern cataract surgery also requires other anterior segment biometric measures in its 
determination of intraocular lens power. These anterior segment biometric measures include anterior chamber 
depth and corneal diameter and are used in conjunction with corneal power (curvature) to determine the 
estimated lens position (ELP). 
 
Tomographic techniques such as OCT image the 
cornea in whole (anterior surface, posterior surface, and 
intervening stroma), and hence OCT requires fewer 
assumptions than current diagnostics for measuring 
corneal power [35]. However, to utilize anterior eye OCT 
images for measuring the corneal refractive power, the 
acquired OCT images need to be faithful representations 
of the imaged eye. As a scanned light technique, OCT 
light is affected by transit through the imaging system, by 
transit through the imaged eye, and by motion of the 
imaged eye during scanning. These corrupting artifacts 
need to be addressed and removed from the images to 
create faithful representations of the anterior eye. Once 
the images are corrected, algorithms to recover the 
pertinent clinical values from the imaged structure can be 
used. 
 
Time-domain OCT (Zeiss Visante) was one of the 
first OCT devices used for the purpose of anterior 
segment biometry [36]. However, motion artifacts due to 
its slow acquisition rate reduced the reliability of its 
measurements [7]. With the substantial speed and 
resolution increase from spectral-domain OCT, accurate spherical corneal power measurements have been 
shown [34,36]. Compared to clinical standards, though, the P.I. has shown that the standard deviation of 
SDOCT measures was higher than clinical resolution [37]. With dedicated motion correction techniques, this 
variability decreased to clinical resolution [32], suggesting the speed of conventional SDOCT alone was not 
sufficient to mitigate motion artifacts. With motion correction, we were also able to measure true OCT based 
corneal topography [38]. This global corneal surface measurement has not been available with clinically 
available OCT systems likely due in part to motion artifacts. Also, in contrast to the slower Visante, the limited 
depth range of SDOCT does not allow the measurement of anterior segment biometry as used in OCT based 
intraocular lens selection [36].      
 
The objective of this aim is two-fold. The first objective is to develop and validate algorithms for the 
accurate measurement of anterior segment biometry from OCT images including clinical corneal spherical and 
astigmatic powers and anterior segment dimensions (such as anterior chamber depth and width). The second 
objective is to test the hypothesis that improved anterior segment biometry measurement with OCT will 
improve cataract surgery outcomes in patients with prior LASIK. 
 
The rationale for this aim is that successful completion of this aim will produce OCT images of the 
anterior eye faithful enough for accurate anterior segment biometry. In addition to the clinical rationale, this aim 
also serves to validate anterior segment measures for the posterior segment aim. Our expectation is that the 
more accurate measurement of anterior segment biometry will remove uncertainty in optical characterization of 
the post-LASIK eye and improve cataract surgery outcomes in this population. 
SA2. Justification & Feasibility. Utilizing current commercially available spectral domain OCT (SDOCT) 
systems, the PI has led development and implementation of a suite of technologies that address automated 
segmentation, motion and optical artifact removal, and measurement of pertinent clinical parameters from OCT 
images of the cornea [20,32,37,39]. In these works, we have shown that we can accurately measure surface 
curvatures of manufactured phantoms with less than 10 µm difference from the nominal manufacturer values 

and with better than 10 µm repeatability. In clinical work, we have shown in normal, non-surgically altered 
corneas that we can measure the corneal refractive power with SDOCT comparable to existing clinical 
standards (Fig. 6). The mean difference of the SDOCT measured corneal power against the clinical standards 
was less than 0.10 D. The standard deviation with motion correction was only 0.25 D (14 eyes). Repeatability 
of measurements was under 0.12 D. For clinical context, glasses prescriptions are measured at intervals of 
0.25 D. 
We have performed a preliminary clinical study demonstrating the ability of SDOCT to measure the 
corneal power change due to LASIK [40]. In LASIK, the excimer laser reshapes the anterior cornea to change 
the eye’s refraction. Hence, the measured refractive change in an eye before and after LASIK should ideally 
equate to the change in measured corneal curvature and power. To test this, 39 eyes of 22 patients were 
enrolled in a Duke University Institutional Review Board (IRB) approved study before they underwent LASIK. 
The enrolled subjects underwent preoperative imaging with topography, Scheimpflug photography, and 
SDOCT and received a manifest refraction (MRx) prior to LASIK. The same measurements were then 
performed at least 3 months after LASIK when visual acuity stabilized. Table 1 summarizes the results. 
 
MRx – Topography 
MRx – Scheimpflug 
MRx –  OCT 
ΔΦK (D) 
-0.50 ± 0.68 
p = <0.001  
ICC = 0.92 (0.85-0.96) 
0.06 ± 0.85 
p = 0.714 
ICC = 0.94 (0.88-0.97) 
0.13 ± 0.60 
p = 0.216 
ICC = 0.96 (0.93-0.98) 
|ΔΦK| (D) 
0.72 ± 0.45 
0.53 ± 0.66 
0.48 ± 0.37 
Table 1. Imaged corneal power change from LASIK compared to vertexed manifest refraction (MRx) change 
In this LASIK population, the mean absolute difference (colored row in Table 1) is the pertinent metric as it 
represents any deviation, positive or negative, from the reference MRx measure. As a point of reference, the 
step size of intraocular lenses is 0.50 D. Topography was found to be significantly different from the reference 
manifest technique (p < 0.001, t-test with generalized estimating equations). On average, the SDOCT 
measurement deviated least from the refractive change and was the least variable. The variability of the 
Pentacam (Scheimpflug) measurement is unacceptably high; as a result of other studies (not this one), 
Pentacam is no longer on the American Society of Cataract and Refractive Surgeons’ clinically accepted 
modality list (iol.ascrs.org). 
 
In addition to spherical measures, we were also able to use OCT to characterize changes in 
astigmatism as seen in Figure 7. In our population, the mean paired difference between the surgically induced 
change in manifest astigmatism and OCT was 0.49 D ± 0.37 D. This was lower than that found with 
topography simulated keratometry (0.60 D ± 0.51 D) which was found to be significantly different from the 
reference manifest technique (p = 0.003, t-test with generalized estimating equations). 
 
Overall, these preliminary results show that our OCT based corneal reconstructions are sufficiently 
accurate to measure the corneal power change both spherically and astigmatically resulting from LASIK.  
SA2. Research Design. Algorithm development and validation. In contrast to the SDOCT system that we used 
for our preliminary data above, the proposed whole eye SSOCT system will provide deeper and wider anterior 
eye views, while simultaneously imaging the posterior eye. There are multiple benefits to this. First, imaging 
more of the anterior eye will allow for complete anterior segment biometry measurements which can be used to 
improve intraocular lens selection [41-45]; the cornea alone is not the only determinant of intraocular lens 
power. Secondly, the co-registered posterior eye image from whole eye OCT will also allow for definition of the 
visual axis to provide a functional reference axis for the corneal power measurement. 
With the raw anterior eye OCT images, the following post-processing algorithms will be applied: 
segmentation of ocular surfaces, removal of optical and motion artifacts, and finally recovery of the desired 
clinical parameters. For segmentation, we will use our prior dynamic programming based algorithms [39] and 
modify them to account for the wider FOV in the whole eye OCT system. In addition, these intensity and graph 
based algorithms will be extended to segment the iris surface to identify the pupil location. 
For removal of optical artifacts, we will directly apply our prior, validated three-dimensional refraction 
correction algorithms [20]. Successful 3D refraction correction requires accurate characterization of sample 
arm scanning geometry, sample space, and refractive indices. The scanning geometry is known from our 
Zemax design, and sample space will be calibrated as described in Aim 1 and [20]. Refractive indices for OCT 
based refraction correction requires knowledge of the group (not phase) refractive index at the wavelength of 
interest. Co-Investigator Dr. Izatt, in collaboration with OCT inventor Dr. David Huang, first described physical 
measurement of the group refractive index of samples (such as the cornea) using OCT in 2004 [46]. Briefly, a 
cuvette of known, calibrated internal chamber thickness is imaged and measured first. Then a representative 
sample (for example, calibration phantom or fresh ex vivo cornea) is placed in the chamber and the thickness 
is re-measured. The image provides the resultant optical pathlength change of the chamber as a result of the 

 
 FIGURE 7. Topography and OCT based total power & 
topography maps. This figure shows that our recovered 
OCT based surfaces are sufficiently accurate to represent 
astigmatic information even in corneas altered by LASIK. 
Before LASIK (top row), topography and OCT are 
comparable (A & B). After LASIK (bottom row), topography 
(D) shows an effectively spherical cornea because it is only 
measuring the ablated anterior surface. In the OCT map (E), 
residual corneal astigmatism is still present (lighter blue) 
likely due to the unchanged astigmatic posterior surface (C & 
F). The measured spherical change pre to post LASIK was: 
ΔMRx = 5.25 D, Δtopo = 3.94 D, ΔOCT = 5.63 D. The 
surgically induced change in astigmatism was: ΔMRx = 2.80 
D, Δtopo = 2.04 D, ΔOCT = 2.63 D. These maps show the 
central 5 mm of the cornea. 
sample for that OCT light. Because the empty physical 
thickness of the chamber has been measured, all 
quantities necessary to determine the group refractive 
index for the intervening representative sample at that 
OCT light bandwidth are known. We have used and 
validated this technique in our prior work with λ0 = 840 
nm [20,37]. We will do the same for the cornea using 
averaged ex vivo samples with the 1060 nm swept 
source light used in this project. 
For motion correction, our prior distributed 
scanning 
OCT 
technique 
used 
both 
increased 
acquisition speed and motion estimation to improve 
measurement of corneal power over standard linear 
scanning SDOCT [32]. Our distributed scanning 
acquired an individual corneal profile in 1/5th the time 
required for a standard linear scan, and after motion 
estimation, still produced a B-scan of comparable A-
scan density to a normal linear scan. The whole eye 
OCT system operates at an A-scan rate of at least 100 
kHz compared to 20 kHz for SDOCT. By applying 
distributed scanning OCT to our whole eye OCT 
system, we can achieve another 5-fold decrease in 
profile acquisition time which should further mitigate 
any 
potential 
motion 
artifacts. 
Additionally, 
the 
presence of the iris in the whole eye OCT images 
offers additional image registration targets that can be 
used for improved motion estimation if needed. 
  
Finally, the corneal curvatures and refractive powers (both spherical and astigmatic) will be recovered 
using our previously described algorithms [37,38]. Importantly, we will be able to align our measurements to 
the reference axis from the entrance pupil to the fovea due to the addition of the iris and posterior eye 
information from whole eye OCT. This provides an advantage over standard keratometry or standalone corneal 
OCT which are aligned along an arbitrary and non-functional axis. 
 
We will confirm the accuracy of these anterior eye algorithms for segmentation, artifact removal, and 
motion correction combined with the whole eye OCT system by using calibration phantoms (glass spheres and 
contact lenses of known curvature and refractive power) and in an initial small series of 20 normal subjects. 
Clinical study. We will enroll subjects who have had previous laser refractive surgery (LASIK, PRK) and 
who will be undergoing cataract surgery to test the hypothesis that improved corneal power measurement 
using OCT will improve the refractive outcomes compared to standard of care (Fig. 8). Each enrolled subject 
must be able to have optical biometry measurements (IOLMaster, Carl Zeiss Meditec) needed for standard of 
care IOL selection. The optical biometric measurements include corneal power (K), anterior chamber depth 
(ACD), and axial length (AL) among others. These subjects will also be imaged with the whole eye OCT 
system. We will then choose two intraocular lens powers predicted to achieve a near emmetropic (plano) 
outcome: 
1. Standard of care (SoC) based on the Haigis-L output recommendation at http://iol.ascrs.org. This IOL 
calculator website for post refractive surgery eyes is maintained by the American Society of Cataract and 
Refractive Surgeons (ASCRS) and is readily available to the clinical community. The website has multiple 
formulas but only two – the Haigis-L and the Shammas-PL – do not require prior records. This is most 
comparable to an OCT based method which also would not require prior records. We are choosing the Haigis-
L because there are prior results in the literature from a small pilot study directly comparing the Haigis-L to a 
different corneal SDOCT based method [45] which can be used to power our study. 
2. Whole eye OCT based. We will use the OCT based vergence tracing formula described in [36]. 
Current clinical IOL formulas have many assumptions inherent in them such as a non-surgically altered cornea 
and that estimated lens position is a function of the non-surgically altered cornea [47]. To remove these 
assumptions which are violated in post LASIK patients, we will recover the pertinent biometry measures 
directly from whole eye OCT and use the described vergence tracing formula to determine the whole eye OCT 
recommended IOL. This formula was developed specifically for OCT, but available time domain OCT at that 

 
FIGURE 8. Overview of primary clinical outcome to determine if 
more accurate anterior segment biometry measurement from 
OCT improves post-LASIK cataract surgery refractive 
outcomes. Before cataract surgery (left column), we will calculate 
the needed intraocular lens (IOL) power via (1) standard of care 
(SoC: 21.0 D in this example) and via (2) our OCT based corneal 
power and anterior segment biometry measurement (20.5 D). The 
surgeon will perform cataract surgery and implant an IOL (21.0 D 
IOL). Based on the post-operative refraction, the power actually 
needed for a plano result will be calculated (20.25 D). The 
difference between this calculated reference and the predicted 
power is the prediction error. We will compare the proportion of SoC 
and OCT results with prediction errors ≤ 0.50 D. 
time was unreliable due to motion artifacts from 
slow acquisition [36]. Current clinical SDOCT is 
faster, but its limited depth range prevents 
measurement of the pertinent biometry values 
used in this formula. Our whole eye system is 
faster than SDOCT and has the depth range 
necessary to measure all the pertinent biometry 
values required to take advantage of this formula.  
After undergoing cataract surgery with an 
intra-capsular posterior chamber IOL of the 
surgeon’s choice, we will measure the manifest 
refraction at the final, post-operative month 1 visit. 
From this post-operative manifest refraction, the 
IOL power required to have achieved an 
emmetropic (plano) result will be calculated [45]. 
We will then compare this calculated reference 
power to the predicted power from the SoC and 
OCT methods and determine a prediction error for 
each. Given that IOLs are only powered in 
discrete increments of 0.50 D, a “good” clinical 
outcome should be within 0.50 D of the target. 
Hence, we will compare the proportion of subjects 
in the two groups (SoC and OCT) who have 
prediction errors within ± 0.50 D of the calculated reference. Though finer discrimination than 0.50 D would be 
desirable, we are limited by the resolution of the IOL at ± 0.50 D steps. We cannot insert a lens that is 20.13 D, 
for example, if that is the theoretical suggested power. Reporting results within ± 0.50 D is also a standard 
accepted and used by the refractive surgical community [48]. 
 
From the literature, we can estimate the expected proportions for the SoC (PSoC) and OCT (POCT) 
groups. Using the Haigis-L, the proportion of subjects who had a result within ± 0.50 D of the calculated 
reference was approximately PSoC = 41.98% [45,49]. In a small pilot using averaged 2D corneal SDOCT 
measurements with less sampling than our 3D reconstructions, Tang et al. found the proportion of subjects 
with a result within ± 0.50 D of the calculated reference to be POCT = 68.75%. Our null hypothesis is OCT will 
equal SoC (H0: PSoC – POCT = 0), and the alternate hypothesis is that OCT will be different (HA: PSoC – POCT ≠ 0). 
Proportions will be compared using a conditional McNemar’s test for paired proportions. Using the estimated 
proportions, α = 0.05, and assuming zero correlation between pairs, we will need a minimum of 60 subjects to 
complete the study to have 80% power to detect a significant difference between these two proportions (PROC 
POWER, SAS v.9.3). If the correlation between pairs is 0.2, 60 subjects will provide 88% power; if the 
correlation is as high as 0.4, the power will be 96%. To account for possible enrollment losses from intra-
surgical complications or follow-up losses, we will target an enrollment of 65 subjects over the project period to 
ensure that we will meet the minimum of 60 completed study subjects. 
In addition to the primary outcome measure above, we will additionally report the mean absolute 
prediction error for the two groups. We will also repeat the optical biometry and OCT imaging post-operatively 
to further study the relationship between pre-operative biometry and the final IOL position (ELP). These 
measurements will allow us to further characterize the relationship between clinical optical biometry and the 
whole eye OCT pre- and post-operatively. 
SA2. Expected Outcomes. We will have developed algorithms to accurately measure anterior segment 
biometry and corneal power with the whole eye OCT system. We will further have clinical data to support the 
idea that accurate measurement of anterior segment biometry will improve cataract surgery refractive 
outcomes in those who have previously had laser refractive surgery. While this growing population would 
particularly benefit because of the unique challenges presented by their altered corneas, the developed 
algorithms can also be applied for the larger group of cataract surgeries in those with normal corneas as well. 
Finally this aim also complements the posterior segment aim by independently validating the measurements of 
anterior segment surfaces by the whole eye OCT system.  
SA2. Potential Problems & Alternative Strategies. While we expect that improved anterior segment 
biometry measurement with OCT will improve cataract surgery outcomes in those with prior LASIK, additional 
parameters could potentially further reduce the IOL prediction error. For instance, as described earlier, modern 

FIGURE 9. Representative original and spatially 
(numerically) corrected retinal OCT B scans (top R) and 
comparison to MRI (bottom L and R) [4]. The posterior eye 
appears flat in the original image (top L) and has a more 
plausible physiologic curvature after spatial correction (top R). 
Reported values in the literature for posterior eye curvature (Rc) 
range from 10.8-14.1 mm [34]. Spatial correction in the 10 eyes 
in this study significantly reduced the difference between OCT 
and MRI posterior eye shape (bottom L). In this small preliminary 
study, there remains a bias and a possible trend between the 
OCT measures and MRI (bottom R). 
IOL power formulas estimate the final position of the IOL by using the corneal curvature. Interestingly, in 
recently presented preliminary data [50], a formula which did not utilize corneal curvature for estimating the 
lens position [51] appeared to do better than those that did use corneal curvature in post LASIK patients 
undergoing cataract surgery. In our study, we will be obtaining additional biometric measurements pre and 
post-op – both standard of care and OCT based – to be able to analyze other contributions to the IOL error if 
there is a systemic bias even after corneal power correction. Regression to these collected data and 
measurements could be utilized as an alternative approach [45].  
Alternatively, other commercially available spectral domain OCT could be used for this aim. However, 
they currently lack the imaging range necessary to perform the anterior segment biometry required for this aim, 
and they generally cannot provide a reference axis to the fovea for their corneal measurements. Conversely, 
the corneal power specific portions of our algorithms could be applied to existing OCT to improve their corneal 
power measurements and benefit the clinical OCT community in general. 
Specific Aim 3: Develop and validate algorithms to produce accurate three-dimensional posterior eye 
OCT reconstructions to supplant MRI for measuring posterior eye shape in myopia applications.   
SA3. Introduction. Current clinical OCT images of the posterior eye are spatially inaccurate. This can be 
observed simply by noting that almost all clinical OCT images of the retina appear artificially flattened when the 
expectation is that the eye globe should be curved (Fig. 9). Because of this spatial inaccuracy, OCT cannot in 
its current form replace MRI for the quantitative study of ocular shape deviations in diseases such as myopia, 
even though OCT is more clinically accessible and has higher spatial resolution than MRI. Further, the spatial 
inaccuracy with OCT is compounded by variable optical artifacts from the system and eye, meaning that image 
differences between patients or time points may not be solely due to pathology as is usually assumed. 
 
The spatial inaccuracy of current posterior 
eye OCT images results from the current standard 
practice of simply displaying A-scans next to each 
other in a parallel fashion. This fails to account for 
the actual, non-parallel transit of OCT light through 
the eye. The objective of this aim is two-fold. The 
first objective is to develop and validate algorithms 
for accurately portraying the posterior eye in OCT 
images by ray tracing the OCT imaging paths 
through the imaged anterior eye. The second 
objective is to then test the hypothesis that 
corrected posterior eye OCT images will be 
comparable to MRI of the same eyes.  
 
The rationale for this aim is that successful 
completion of this aim will produce posterior eye 
OCT images faithful enough to replace MRI 
imaging 
for 
studying 
posterior 
eye 
shape 
biomarkers. 
Our 
expectation 
is 
that 
this 
development will remove cost and logistical 
barriers limiting more widespread study of posterior 
eye shape as a diagnostic biomarker for myopia 
progression, particularly in children and young 
adults. 
SA3. Justification & Feasibility. In an initial study 
by the P.I. [4], we enrolled subjects in an IRB 
approved study who had recently obtained head 
MRIs for their medical care but did not have ocular 
disease. We performed ocular biometry and retinal 
SDOCT on these subjects. Using commercial ray tracing software, we modeled the commercial SDOCT 
system and the subjects’ eyes using their measured biometric data. The spatial orientation of the OCT scan 
paths was then used to reorient each OCT A-scan to its modeled location in MATLAB. Using this correction, 
there was a significant reduction in differences and variability between posterior eye shape measured by OCT 
and MRI (Fig. 9). 
Despite the overall improvement in posterior eye OCT shape in this pilot study, there remained a bias 
between the OCT measures and MRI. The possible sources of error in this preliminary work (other than small 

FIGURE 10. General overview of clinical study design to 
determine the difference between OCT and MRI posterior 
eye images. Using radius of curvature as a general measure of 
shape, we will compare the paired differences between the two 
modalities. In a second study in subjects with staphylomas, we 
will use root mean square difference to describe the topographic 
differences between the two registered modalities.  
sample size) include the lower resolution of these non-orbital MRI scans, the use of schematic model anterior 
eyes, motion errors, and the unknown system-to-cornea distance to control for scan magnification. 
In this proposal, we will perform a prospective study with dedicated orbital MRIs, with measured and 
model-free determination of all ocular surfaces, with decreased patient motion from the faster SSOCT system, 
and with direct measurement of the system-to-cornea distance (see SA1.Research Design.System 
Calibration). With these improvements, we anticipate that we will decrease the curvature difference between 
MRI and OCT to less than a millimeter (resolution of our MRI). Further, we will compare surface topographies 
of myopic patients with known morphometric changes (e.g. staphylomas) imaged with MRI and OCT.  
SA3. Research Design. Algorithm development and validation. We will extend our previously developed 
three-dimensional refraction correction algorithms [20] to perform ray tracing of the OCT scan paths through 
the ocular surfaces contained within the whole eye OCT images. The anterior eye image content will be 
segmented and corrected of optical distortions as in Aim 2 (refraction, non-telecentricity, motion). Group 
refractive indices for our light source (λ0 = 1060 nm) will be used for refraction correction as described in detail 
in SA2.Research Design.Algorithm Development. For the lens, we will measure the group index for IOLs, and 
use a polynomial model for natural lenses [52] as was used in our pilot work [4]. The retinal layers of the 
posterior eye will also be segmented in preparation for post-processing [53]. With a known system topology, 
measured and hence known system-to-cornea distance per patient, and an accurate representation of the 
anterior eye components, we will determine the spatial orientation of the OCT scan paths to the posterior eye. 
With these OCT scan paths, the posterior eye A-scans will be properly repositioned spatially to restore 
anatomic accuracy.  
 
We will validate the algorithms by imaging a constructed model eye. The model consists of a custom-
made +30 D rigid gas permeable contact lens (Essilor, center thickness 490 µm), a +20 D intraocular lens 
(SA60AT, Alcon), and a glass hemisphere with a 12.5 mm radius of curvature (64-496, Edmund Optics). These 
will be mounted into a lens tube at physiological distances and then immersed into balanced salt solution, 
leaving only the anterior surface of the contact lens (“cornea”) exposed. The model eye will be imaged with the 
whole eye OCT and the surfaces recovered. After correction with the above algorithms, the curvature of the 
hemisphere will be compared to that of the 
manufacturer reference. 
Clinical study. For the clinical studies, we will enroll 
two groups of subjects: 1. pseudophakic subjects 
with no known history of ocular pathology other 
than refractive error (no more than -6 D) and 2. 
pseudophakic and phakic subjects with high 
progressive 
myopia 
including 
those 
with 
staphylomas. 
Group 1. The goal for the first group is to 
demonstrate in a normal eye the ability to 
accurately recreate the posterior eye with OCT 
spatially comparable to MRI (Fig. 10). We will only 
include monofocal pseudophakic individuals in this 
group. This reduces uncertainties by using a known manufactured lens. At the Duke Eye Center, the 
monofocal lens of choice is manufactured by Alcon, and the company has provided us with the optical 
prescriptions of their IOLs for research use under a non-disclosure agreement. With the optical properties of 
the IOL known, we anticipate being able to fully characterize these normal, pseudophakic eyes optically. This 
also provides a baseline from which phakic eyes can be studied.  
Each enrolled subject will undergo the following imaging: optical biometry (LenStar, Haag-Streit), whole 
eye OCT, and MRI imaging. MRI will be performed on the clinical research scanners at the Duke Brain Imaging 
and Analysis Center (see support letter from director Dr. Allen Song). Annually, with the assistance of 
dedicated research coordinators, they image over 2000 subjects with MRI for over 70 research projects. Their 
scanner (GE) has a magnetic field strength of 3.0 Tesla. A high resolution T1 FSPGR sequence will be used to 
acquire images of the subject's eyes. Images will be acquired in the axial plane, with FOV=25.6cm, TR=8ms, 
TE=3ms, flip angle=12, 512x512x50 imaging matrix, completely covering the eyes. A second T2 weighted 
sequence will be acquired at the same slice prescription with TR=3s and TE=25ms. With these scan 
parameters, each voxel will be 0.5 mm3. During scanning, a fixation target will be provided to minimize eye 
motion during acquisition. 

 
The radius of curvature of the posterior eye will be measured via surface fitting of the posterior eye data 
from OCT and MRI. To have 80% power to detect a 0.5 mm difference in radius of curvature between OCT 
and MRI using a paired one-sample t-test, we will need to enroll a minimum of 48 normal pseudophakic 
subjects (PROC POWER, SAS v9.3). This power calculation assumes that with our whole eye OCT and 
improved algorithms, we will halve the standard deviation of the difference  between MRI and OCT found in our 
initial study (2.3 mm), resulting in an estimated standard deviation of 1.2 mm. At a 0.5 mm difference, we will 
be at the effective voxel resolution of the MRI scans and should also have more than enough resolution to 
identify the large-scale topographic abnormalities noted in previous MRI studies of myopia [25,26]. 
Group 2. After enrollment and data analysis of the first group, we will then examine the ability of OCT to 
provide spatially comparable images to MRI in eyes that are structurally abnormal. We will enroll 20 adult 
patients with progressive high myopia (more than -6 D) especially targeting those with posterior staphylomas – 
outpouchings of the posterior eye. These subjects can be pseudophakic or phakic. Phakic subjects will be 
required to be cyclopleged to ensure no dynamic lens changes during imaging. For phakic reconstructions, we 
will use age appropriate models of the natural lens indices as we have used in our previous study [52]. After 
recovery of the OCT and MRI surfaces, the OCT surface will be registered to the MRI surface at a common 
feature (e.g. optic nerve) and via cross-correlation. A root mean square deviation across the registered 
surfaces will then be calculated to quantify the topographic difference between the OCT and MRI surfaces. 
SA3. Expected Outcomes. At the conclusion of this aim, we will have developed algorithms to spatially 
correct full eye OCT images and validated them with reference phantoms. We will also have tested the idea 
that spatially corrected OCT images will be comparable to MRI representations of the posterior eye first in 
pseudophakic normal eyes and then in psuedophakic and phakic myopic eyes. This will provide the basis for 
OCT to replace MRI for ophthalmic studies of ocular morphometry, lowering cost and logistical barriers to 
these types of studies while also providing higher resolution images over MRI. This has clinical utility for study 
of myopia progression and risk stratification based on prior studies associating deviations in ocular shape with 
pathologic myopia. 
SA3. Potential Problems & Alternative Strategies. The FOV of our proposed posterior eye OCT channel will 
be 85° (or ~120° relative to center of the eye) – at least double that of standard retinal photography. This 
should be sufficient for measuring posterior eye shape as it encompasses the entire macula and reaches 
nearly to the equator of the eye. Should larger FOV be desired, though, we can use montaging as others have 
shown with OCT [54] to further increase our FOV. Because our whole eye OCT also concurrently images the 
anterior eye, our reconstructions can account for the relative change in position between the system and the 
eye to increase the accuracy of montaging and provide additional references for montage reassembly. 
While the natural lens does present an ambiguity, by first imaging pseudophakic subjects in group 1 
and part of group 2, we will first refine and validate our algorithms independent of the natural lens. This lays the 
foundation for subsequently investigating and isolating any differences between MRI and OCT resulting from 
imaging phakic subjects in group 2. Our preliminary algorithm using a polynomial gradient index lens model 
[52] demonstrated reduced differences between OCT and MRI [4]. Alternate OCT based natural lens models 
[55] can be incorporated into our reconstructions for reanalysis of captured and stored image data if necessary. 
Our use of these gradient indices is far more sophisticated than the average refractive indices successfully 
used for ocular axial biometry both in research and clinically [56,57 and commercially available partial 
coherence interferometers]. If necessary, we can similarly use average indices. Importantly, with the validated 
pseudophakic group, we will be able to compare and contrast quantitative performance of OCT in phakic 
subjects. This data has important implications 
for both the OCT and myopia communities.   
Timetable. We anticipate this ambitious 
project will require the full requested five year 
project period to complete. We anticipate the 
whole eye SSOCT system design and 
assembly will take approximately one year. 
Algorithm development will begin with initial 
images of phantoms and normals and can 
occur 
concurrently 
with 
clinical 
imaging 
because all acquired images can be post-
processed at a later date if necessary. Data 
collection, data analysis, and publication will 
occur on a rolling basis throughout the project.