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Optical coherence tomography (OCT) imaging has become widespread in ophthalmology over the past 15 years, because of its ability to visualize ocular structures at high resolution. This article reviews the history of OCT imaging of the eye, its current status, and the laboratory work that is driving the future of the technology.
Optical coherence tomography (OCT) has advanced considerably since it was first applied to the eye.1–7 It is an extension of a technique called low-coherence interferometry, which was initially applied to ophthalmology for in vivo measurements of eye axial length.1 At the time of introduction, it was used to obtain in vivo optical cross sections of the anterior segment,6 as well as retinal diseases, such as macular detachment, macular hole, epiretinal membrane, macular edema, and idiopathic central serous chorioretinopathy.4 OCT cross sections were also used to evaluate the optic disc and retinal layers5,8 such as the retinal nerve fiber layer (RNFL).9 Scan patterns that enabled reproducible measurements were developed,10 and these eventually became incorporated into a commercial system, which had an axial resolution of ~10 μm.
The first clinical system was limited to a scanning speed of 400 axial scans (A-scans)/s because of a physical constraint: a moving reference mirror. OCT uses low coherence interferometry to obtain A-scan intensity profiles, and the process requires light to be split and sent to both a reference arm with a mirror and to the sample. Provided the path length to the reference mirror and tissue match to within the coherence length of the light source, when the reflected beams recombine, interference occurs. Intensity information, in the form of a reflectivity profile in depth, can be extracted from the interference profile. Changing the location of the reference mirror allows backscattered tissue intensity levels to be detected from different depths in the tissue sample. This approach is referred to as time-domain (TD)-OCT because time-encoded signals are obtained directly. Several improvements in OCT hardware have been introduced since the first commercial TD-OCT system became available. Better axial resolution11–13 and increased scanning speed14–23 are the two main advancements that have recently become incorporated into commercial systems.
The implementation of broadband light sources into OCT systems11 improved the axial resolution from ~10 μm to as high as 2 μm in tissue.24 Acquisition speed has improved considerably by detecting backscattering signals in the frequency domain,14–23 which means backscattered depth information at a given location can be collected without the movement of a reference mirror. Frequency information is acquired with a broad-bandwidth light source, charge-coupled device (CCD) camera, and a spectrometer14,17,18,20 or by sweeping a narrow-bandwidth source through a broad range of frequencies with a photodetector.16,21–23 The approach that incorporates a broadband light source is often referred to as spectral-domain (SD)-OCT, whereas the latter is termed swept-source (SS)-OCT. In both approaches, intensity profiles (A-scans) are obtained using a Fourier-transform of the detected frequencies, and this facilitates rapid A-scan collection. In addition to improved scanning speed, frequency-domain-OCT also offers the advantage of higher detection sensitivity—that is, it exhibits higher signal-to-noise, given a perfect reflector.23,25 A summary of OCT detection techniques can be seen in Table 1.
With these speed and sensitivity improvements, it is now feasible to collect volumetric (three-dimensional; 3D) scans of tissue, whereas in the past, the amount of time required to do this would have been prohibitive. Broadband volumetric retinal imaging with SD-OCT at speeds of up to 312,500 A-scans/s26 and SS-OCT at 249,000 A-scans/s27 have been demonstrated. To date, most clinical systems operate at an acquisition rate of ~27,000 A-scans/s and an axial resolution of 5 to 6 μm. A summary of the current commercially available retinal imaging systems is presented in Table 2.
Conventionally, the most common scan patterns in TD-OCT glaucoma imaging were a 3.4 mm scan around the optic nerve head (ONH) and six equally spaced radial scans through the macula (6 mm) and optic nerve (4 mm). RNFL thickness is obtained via automated RNFL segmentation in the circumpapillary scan protocol, whereas macular thickness (internal limiting membrane [ILM] to the photoreceptor inner segment-outer [IS–OS] segment junction) is automatically segmented in the macular scan pattern. The optic nerve scan is used to obtain cup area, disc area, cup diameter, disc diameter, and rim area. These ONH parameters are obtained automatically: the software detects the ONH margin/RPE tips, but the user can modify the location if the ONH margin detection algorithm is inaccurate.
The ONH and RNFL scan protocols have been used since TD-OCT became commercially available, and RNFL and ONH parameters have been shown to differ between glaucomatous and healthy eyes.9,28–33 The glaucoma-discriminating ability, measured by the area under receiver operating characteristic curves (AROC) of RNFL (AROC = 0.94) and disc parameters (e.g., rim area AROC = 0.97), has been reported to be higher than macular volume and thickness (AROC, both 0.80).34 A similar glaucoma-discriminating ability is seen in comparing TD-OCT imaging and SD-OCT imaging when similar parameters are examined.35 However, it may be possible to further improve glaucoma discrimination using parameters obtained from 3D scanning. With the commercialization of rapidly scanning SD-OCT systems, 3D volumes of tissue are now easily acquired. A 3D dataset not only allows a quantitative analysis from more locations but, once a volume has been collected, OCT fundus (en face) images can be obtained by integrating A-scans.20 These can be used for a subjective assessment of signal quality and to assist with evaluating and/or correcting eye motion that may have occurred throughout the scan. The OCT fundus image also allows registration of OCT cross sections to precise retinal locations.
Acquisition of 3D datasets has led to the advancement of software methods for efficiently analyzing and summarizing these vast amounts of data. One method of obtaining RNFL thickness measurements has been sampling the 3D volume (e.g., 6.0 × 6.0 × 2.0 mm, centered on the ONH) after acquisition, at a diameter of 3.4 mm centered on the ONH (Fig. 1). This method has been shown to have higher reproducibility than the conventional TD-OCT 3.4 mm scan circle, where the image is acquired along the circle only.36 One explanation for the improved performance is that, with TD-OCT, scan placement is dependent on the user and can be variable, but with SD-OCT, the circle can be consistently placed in the same location by using landmarks within the 3D volume.
Although sampling 3D volumes after acquisition may be an effective way of summarizing RNFL measurements, it is doing so at a cost: data outside the 3.4-mm band are not being used. Subjectively, wedge defects and global thinning may be easy to spot, but subtle changes or deviations from normal outside the 3.4-mm sampling band may be missed. One way of addressing this is to create an RNFL thickness map, which consists of all thickness measurements outside of the ONH. From this, thickness measurements from one subject can be compared to population thickness measurements. To date, however, commercial software is available for looking at deviation from normal, but there is no quantitative assessment using all available RNFL information.
Different approaches have been proposed for quantifying 3D data. 3D RNFL thickness has been analyzed in terms of a thickness profile as distance from the ONH increased.37 In healthy eyes, the slope of RNFL thickness increases near the margin of the ONH, peaks, and then decreases with increasing distance from the ONH center in all but the nasal quadrant, which linearly decreases starting from the disc margin. Another approach (Ishikawa H et al. IOVS 2009;50:ARVO E-Abstract 3328) exploits 3D macular data, which have been summarized using segmentation of the inner retinal complex (IRC: retinal ganglion cell layer [RGC], inner plexiform layer [IPL], inner nuclear layer [INL]; Fig. 2).38 This approach reduces IRC data to superpixels (4 × 4 adjacent sampling points) and compares these superpixels to a normative thickness superpixel dataset. By condensing measurements into superpixels, it is less likely that small imaging artifacts or algorithm failure will have an effect.
One commercially available system has developed an approach for summarizing macular data called the Ganglion Cell Complex (GCC; RTVue, Optovue, Inc) which consists of essentially the same layers as the IRC: the RGC (retinal ganglion cell bodies), RNFL (RGC axons) and IPL (RGC dendrites). The GCC measurements are then directly compared to a normative database and thickness difference and significance maps are available (Fig. 3).
While a comparison to a normal population may reveal differences, structural changes may be occurring while a patient remains within normal limits and therefore go undetected. Ideally, a longitudinal comparison could be made for a given individual to look for subtle structural changes attributed to disease progression. One approach, proposed by Kim et al.,39 is to allow compatibility between TD-OCT and 3D OCT device iterations. Since TD-OCT devices have been commercially available longer than 3D imaging systems, years of patient information may be available. The method presented by Kim et al. resamples a 3D-OCT dataset for every possible 3.4-mm circular scan location within the boundaries of the 3D-OCT volume. It then uses cross correlation between these virtual circular scans and the TD-OCT 3.4-mm scan to automatically match the TD-OCT scan circle location within the volume.
To longitudinally compare 3D volumes, however, image registration techniques must be developed to spatially align 3D scans before they can be compared. This may be accomplished using cross-correlation,40,41 or by using landmarks within the OCT fundus image, such as blood vessels.42,43 Eye motion during acquisition has been shown to alter scan location,44–46 and the effect of eye motion is visible on OCT en face images as discontinuous blood vessels. Detecting and correcting blood vessel location to align the OCT fundus can help correct eye motion,42 which may be useful for cross-sectional analysis.
Longer wavelength imaging (~1-μm center wavelength) of the lamina cribrosa27 and birefringence imaging of the RNFL using polarization-sensitive (PS)-OCT47–49 are two techniques under development that may improve the diagnosis and monitoring of glaucoma. These are further described in the Preclinical and Laboratory Studies section of this review.
The macular scan pattern discussed above—six radial macular scans, 6 mm long, spaced 30° apart—has traditionally been used in TD-OCT imaging to assess retinal parameters such as total retinal thickness and the IRC. Three-dimensional imaging; however, has revolutionized the examination of retinal disease.50–55 An examination of the 3D structure of the retina, as opposed to just six radial scans, may make subtle structural changes apparent. For example, using high-resolution 3D imaging to observe the photoreceptor IS–OS junctions may be an indicator of visual outcomes after macular hole surgery.56–58
At present, one application of 3D OCT of imaging retinal diseases that has considerable clinical potential is surgical planning and the evaluation of surgical outcomes. The use of OCT for planning an access point to release the hyaloid for vitrectomy using the six radial scan pattern in TD-OCT has been described.59 Although this was effective for minimizing traction forces on the macula during surgery, a detailed 3D map of the hyaloid membrane and subhyaloid space could further inform the clinician. Falker-Radler et al.60 used 3D imaging to visualize the vitreomacular interface in subjects who were undergoing surgery for epiretinal membrane. Others have used OCT for evaluation of structure after surgery for macular hole57,58,61 and vitreomacular traction.52,62–64 The use of 3D imaging for surgical preparation and evaluation of surgical outcomes has the potential to improve with the use of longer wavelength imaging, which is described later in this review. Automated segmentation of structures of interest, when possible, may provide objective measurements to clinicians for pre- and postsurgical evaluation.
Quantification of thickness is possible in certain diseases,65,66 especially with early stage changes.67,68 The reproducibility of SD-OCT retinal thickness measurements is higher in than that of TD-OCT.69 Thickness has been shown to correlate with best corrected visual acuity in diabetic macular edema70 and ERM.71 Although thickness may be a clinically useful correlate of visual function, there are cases and diseases in which no correlation to thickness is seen, and thus clinicians should exercise caution in interpretation of thickness measurements.72
Drusen volume may be a predictor of progression of age-related macular degeneration,65 and efforts are under way for automated assessment.66 Although accurate quantification of volumetric tissue changes will assist with longitudinal monitoring of disease, fully automated segmentation may not be reliable because of shadowing from fluid in the retina73,74 or because of pathologic events that disrupt normal retinal structures, such as macular hole, subretinal fluid, pigment epithelium detachment, and others.75,76
In cases in which fully automated segmentation fails, C-mode visualization of structures may augment subjective analyses.77 A 3D volume of data can be sectioned in any plane after acquisition, and for C-mode visualization, data are sectioned perpendicular to the retina. The section can be of any thickness, so structures embedded within a volume can be exposed. Often, since the true structure of the retina is curved, exact perpendicular sections slice through several layers simultaneously78,79; thus, aligning the volume to structures such as the ILM or RPE assists with isolating structures of interest.77 Figure 4 shows an example C-mode section taken after aligning a macular SD-OCT 3D volume to the RPE. Moving axially, past the retina and RPE, enables visualization of the choroidal blood vessels. The choroidal vessels are not apparent in the corresponding SD-OCT fundus image because of highly reflective layers superficial to the choroid. The C-mode provides alternative viewing perspective for many retinal diseases, such as cystoid macular edema, central serous retinopathy, vitreoretinal traction, and age-related macular degeneration,77 and it can improve the visualization of their pathologic features.
It is also possible to image the choroid by focusing the illuminating OCT beam deeper and moving the choroid closer to 0 delay.80 In addition, longer wavelength imaging at ~1 μm81 allows for deeper penetration of light into the retina and choroid. A combination of these approaches may improve the current understanding of choroidal diseases.
Correcting ocular aberrations with adaptive-optics (AO)-OCT82 may also provide a unique viewing perspective for retinal diseases. This technique has been applied to view photoreceptors83 and RNFL.84 The utility of longer wavelength imaging and AO-OCT is under investigation and is described in the Preclinical and Laboratory Studies section below.
Anterior segment OCT (AS-OCT) provides structural information of the cornea and anterior chamber without contacting the eye, offering an ease of image acquisition and a considerable advantage over ultrasound biomicroscopy (UBM). While it cannot be used to image deep structures such as the ciliary body, as UBM can, AS-OCT has higher axial resolution (5–10 μm for AS-OCT compared with 25 μm for UBM).85
It is possible to acquire high-resolution images of the sclera, angle, and iris with AS-OCT imaging at longer wavelengths (1.3 μm).86 High-resolution images of the anterior chamber angle can also be obtained with 850-nm systems, and this has led to the visualization of the trabecular meshwork and Schlemm's canal.87–89
Raster scanning and radial scanning of the cornea have been used to measure thickness,90 resulting in reliable pachymetric mapping.91,92 Pachymetric measurements obtained with AS-OCT may assist with planning or follow-up of LASIK patients93 or may be used to diagnose keratoconus.94
In addition to its potential benefits in the evaluation of the anterior chamber angle and cornea, AS-OCT has also been shown to be applicable in the assessment of lens thickness in phakic eyes95 or intracorneal ring placement.96 This can provide an alternate, noncontact, method of pre- and postsurgical assessment.
As previously discussed, SS-OCT obtains time-encoded spectral information by sweeping a narrow-bandwidth laser through a broad optical spectrum. Backscattered intensity is detected with a photodetector. This process is in contrast to SD-OCT, which uses a broad bandwidth light source and detects the interference spectra with a CCD camera and spectrometer. The use of spectrometer-based SD-OCT has become widespread in the clinic, but there are some benefits to photodetector-based SS-OCT systems. Similar to SD-OCT, SS-OCT offers speed and sensitivity advantages over TD-OCT.23,27 To date, speeds of up to 249,000A-scans/s have been attained in the eye.27 Therefore, eye motion artifacts are greatly reduced compared with TD-OCT.22
One advantage of SS-OCT over SD-OCT is that it does not require a CCD camera and spectrometer and instead uses a simpler photodetector.27 A drawback to camera-based SD-OCT detection is a drop-off in signal with depth of scanning because of the finite pixel size of the CCD camera.25,97 Although this can be improved by reducing the camera pixel size,97 it increases the complexity and therefore the cost of the CCD array. A noticeable drop-off in signal with depth typically does not occur with SS-OCT imaging due to the narrow bandwidth of the light source.23,97
At this time, one disadvantage of SS-OCT is that most systems are now operating at longer wavelengths (λ = 1–1.3 μm), with very few studies demonstrating SS-OCT in the 800 nm range.98,99 Water absorption limits the usable bandwidth at 1 and 1.3 μm99 and this limits the axial resolution; the water absorption window at 850 nm is larger, so higher axial resolution can be achieved.
While axial resolution at longer wavelengths may not be as fine as at 850 nm, there are advantages to using ~1- and 1.3-μm sources. Posterior segment imaging using ~1-μm (1040–1060-nm)81,100,101 center wavelengths has allowed deeper penetration into the retina, optic nerve head, and choroid,81,102 which may be beneficial for imaging choroidal vessels, lamina cribrosa, and diseases such as choroidal neovascularization.103 The water absorption window at 1.3 μm offers even deeper penetration of light and may be useful for cornea and anterior segment imaging.16,22,23,104–106 Anterior chamber imaging at 1310 nm has been applied to visualize anterior segment structures anterior and posterior to the iris, Schlemm's canal, trabecular meshwork, and the scleral spur,107 as well as the anterior chamber angle.108
While SS-OCT systems at any wavelength are not yet commercially available for clinicians, in part due to the cost of the light source, there is clinical potential for such devices. No signal drop-off with depth in SS-OCT, in combination with deeper penetration from longer wavelengths, may improve delineation of the outer retina, RPE, and choroid thereby enhancing the performance of segmentation algorithms. In addition, high-speed 1.3-μm imaging may expand the use of anterior segment OCT imaging.
Ophthalmic systems that employ adaptive optics (AO) dynamically adjust their optical characteristics to compensate for monochromatic aberrations that occur naturally in the eye. AO was initially proposed109 and later used by astronomers to correct distortions of light passing through the atmosphere.110 In 1997, AO was demonstrated in the eye by Liang et al.,111 who used a Hartmann-Shack wavefront sensor and a deformable mirror to correct contrast sensitivity and improve quality of vision for human subjects and to obtain higher resolution images with an AO fundus camera. Shortly thereafter, individual cone mosaics were imaged.112
AO-OCT was first reported by Miller et al.82 in 2003 to improve transverse resolution. Uncorrected, conventional OCT beams 1 mm in diameter have a transverse resolution limited to ~15 to 20 μm.113 This makes it difficult to visualize individual cellular structures. One way to improve transverse resolution is to increase the numerical aperture, which in practice means increasing the diameter of the OCT beam entering the eye, since this would decrease the spot size on the retina. However, the theoretical diffraction-limited resolution cannot be attained due to ocular aberrations114 that occur when the pupil is dilated.115,116 AO-OCT measures and corrects these aberrations using wavefront sensing and deformable mirrors, thereby minimizing spot size and improving transverse resolution. It should also be noted that aberrations can be dependent on the bandwidth of the light source used for OCT imaging,116 and these may be improved using an achromatizing lens.117
Ultrahigh (axial)-resolution AO-OCT was introduced in 2004, improving transverse resolution to 5 to 10 μm in the retina.113 Zhang et al.118 developed an AO SD-OCT system and saw an enhancement of the photoreceptor IS–OS junction in vivo with AO. C-mode sectioning of 3D datasets have also facilitated the visualization of axon bundles in the RNFL84 and cone photoreceptor mosaics from healthy subjects,83,119 and subjects with structural abnormalities120 and optic neuropathies.121
One disadvantage of AO imaging is that the depth of focus is narrow, which means focusing simultaneously at different depths is difficult. For example, photoreceptors, located deep in the retina, and superficial retinal ganglion cells cannot be brought into focus at the same time. It may be possible to address this limitation by scanning in depth and varying the focal plane122 or by stitching together volumes.123 Another limitation to AO imaging is that the field of view is restricted to approximately 1° to 3°; the use of an eye-tracking system to acquire a series of neighboring scans and gradually build up an image covering a larger volume may provide one solution to this limitation.124
A potential advantage of improved lateral resolution with AO-OCT is improved understanding of normal and pathologic retinal function in vivo. AO may also help to improve the overall quality of images obtained from eyes that have more aberrations. Enhanced lateral resolution and improved image quality may then lead to better performance of automated segmentation algorithms and assist with disease diagnosis and follow-up.
Polarization-sensitive (PS) OCT detects polarization changes in circularly polarized light.125 PS-OCT was initially applied to characterize the birefringence of tooth enamel,126 skin,127 and cartilage.128 In 2001, PS-OCT was first used in the eye129 to measure birefringence of the RNFL in rhesus monkeys. RNFL birefringence was measured in humans by Cense et al.47,48 and Yamanari et al.,49 who found that, unlike RNFL thickness, birefringence does not change as a function of increasing radius from the ONH.48 It does, however, vary by sector around the ONH, with higher birefringence in thicker areas.48 Because birefringence may change with disease, RNFL birefringence obtained with OCT may eventually provide an additional indicator of glaucomatous change. The utility of PS-OCT in glaucoma detection and monitoring is currently under investigation. In addition to measuring RNFL birefringence, a longer wavelength (λ = 1.3 μm) PS-OCT system has been used to observe the anterior chamber in subjects after glaucoma surgery.108 A swept-source PS-OCT system at a 1-μm center wavelength was used to image sclera and lamina cribrosa,130 which may provide insight into structural changes occurring in the ONH in glaucoma.
Polarization of the RPE may be important in the detection of macular disease.131,132 Gotzinger et al.132 developed a segmentation algorithm based on what they refer to as the “polarization scrambling effect ” of the RPE, which provides an alternative to conventional intensity-based quantification. A combined AO PS-OCT system was later used to measure RPE polarization scrambling.133 Conventional PS-OCT was used to observe subjects with AMD,134,135 where abnormal birefringence was co-localized with exudative lesions.
PS-OCT has been used for anterior segment imaging to measure corneal birefringence,136,137 and these measurements were used to compensate for corneal birefringence in retinal imaging.138 A difference in polarization in healthy corneas versus those with keratoconus was demonstrated in vitro, suggesting that PS-OCT may eventually provide insight into corneal diseases in vivo.139
The aforementioned studies indicate that PS-OCT offers an alternative approach for detecting changes of optical properties in tissue. If it can be established that a change in birefringence occurs before tissue thinning or thickening, it may allow earlier detection and the opportunity for earlier intervention.
Subject eye motion can alter the intended location of an OCT scan. Attempts to correct eye motion by using postprocessing methods are under development, but real-time eye-tracking systems may provide an alternate method of avoiding eye motion artifacts.46 Menke et al.140 showed that an SD-OCT system with built-in eye-tracking can provide reproducible measurements, but it is yet to be shown whether this yields higher reproducibility or better sensitivity and specificity than devices without eye-tracking systems.
As described earlier in this review, OCT is already being used for surgical planning and follow-up. In addition, there has been progress in the development of intraoperative OCT systems. Intraoperative OCT was first demonstrated in anterior segment surgery, where a 1310-nm system was coupled to an operating microscope.141 The use of a handheld OCT retinal imaging device has also been demonstrated for use in patients undergoing vitrectomy, after removing either the ILM or epiretinal membrane, to better visualize the macular disease.142 It is possible that the development of an intraoperative approach may be further improved using a projection of a virtual OCT image over the surgical site and within the line of sight of the surgeon,143 but the means of implementing this technique for surgery still has to be investigated.
The noninvasive nature of image acquisition, together with the commercialization of systems optimized for laboratory use has resulted in a recent increase in the number of animal studies using OCT. Two- and three-dimensional scanning with OCT is appealing, because the same animals can be observed over time in vivo, making longitudinal studies of ocular structures possible without the need to kill animals at various time points and obtain histologic sections. Not only does this method reduce the number of animals needed for experiments, it is also superior to cross-sectional experiments that require different animals for different time points. The following briefly summarizes recent studies using OCT in animals, in small to large animal models.
The eyes of small animal models commonly used in developmental biology, such as xenopus laevis larvae144 and zebrafish embryos145 have successfully been imaged with OCT. Rodent imaging with OCT is becoming increasingly popular, given their relatively low cost and short lifespan and therefore shorter time for disease progression. In addition, many transgenic models are easy for researchers to access. OCT has been used to study ocular dimensions146 and characterize normal eye growth147 as well as growth of eyes in mouse models of myopia.148 Mouse models of retinal degeneration have been imaged using TD-OCT,149,150 and healthy and degenerative mice with SD-OCT.151–158
Recently, methods for automatically obtaining measurements from mouse OCT images have been presented.159,160 Images taken in an anesthetized mouse, held in place using a stage with a glass coverslip to neutralize the strong refractive power of the mouse cornea were shown to be reproducible.159 This indicates that 3D SD-OCT imaging of the mouse retina may be useful for longitudinal studies of retinal structure in mice.
Rats also provide an interesting platform for studying structural changes in the retina and optic nerve in response to injury or disease. Given their larger eyes, it is less complicated to focus on the retina than in the mouse. Retina and optic nerve imaging has been demonstrated in rat models of retinal degeneration,151 retinal vein occlusion,161 retinal ganglion cell degeneration post nerve-crush injury162,163 and elevated intraocular pressure,164 suggesting that there is also potential for rats to be used for longitudinal studies with OCT.
The eyes of larger animal models, such as chickens with retinal degeneration,165 have been imaged. Researchers have also used OCT to examine birds of prey,166 pigs,167,168 cats,169 and rabbits.170–178 These animals have eyes that are comparable in size to the human eye, which means large modifications of the OCT system optics are not necessary.
Nonhuman primate models are especially appealing for studies with OCT, since their ocular size and structure closely match those of the human eye. Nonhuman primate imaging may provide novel insight into the mechanical damage to the RNFL and ONH associated with increased IOP, as is often seen in glaucoma. PS-OCT has been used to look at the birefringence of the RNFL,129,179,180 and RNFL thickness was measured in eyes with unilateral, laser-induced ocular hypertension.181,182 Strouthidis et al.183 examined 3D SD-OCT images of the optic nerve in nonhuman primate eyes. They visualized the termination of Bruch's membrane, border tissue, and the anterior scleral canal opening and showed that these structures correlated to disc photos184 and histology.185 This set the stage for a study of alterations in the ONH that are due to increased intraocular pressure.186
Ultimately, longitudinal OCT studies of small and large animals may help evaluate the efficacy of pharmacologic agents, stem cell therapies, surgical intervention, and retinal prosthetics while reducing the required number of animals. OCT has the potential to provide a better understanding of disease development and progression in transgenic and other models of disease, which may eventually translate to improved clinical assessment and understanding of disease.
The use of OCT imaging in ophthalmology has increased steadily in recent years, in part due to technological improvements such as scanning speed, sensitivity and resolution. The field continues to grow and transform the way glaucoma and retinal diseases are monitored. With 3D imaging, there are new ways to visualize pathologic features and corresponding challenges to overcome. Current approaches, such as RNFL thickness maps and C-mode visualization, attempt to summarize structural information. These methods are not necessarily optimized for efficient cross-sectional and longitudinal analysis, but with technological improvements such as longer wavelength imaging, SS-OCT, AO-OCT, and PS-OCT on the horizon, even more structural detail will be available. Translating these techniques to the clinic has already begun and many could eventually be made available to the clinician. A combined slit-lamp/OCT system that allows the clinician to access structural information during a routine examination may eventually be available, as may a portable operating room system. The use of OCT in animal models has the potential to further understanding of disease while offering a platform for testing novel approaches to treatment, as well as for innovations of the OCT technique itself. The future of OCT is promising, but with some constraints that, if history is any indication, will become the basis for future advancement.
Supported in part by National Institutes of Health Grants R01-EY13178 and P30-EY08098; The Eye and Ear Foundation, Pittsburgh, PA; and an unrestricted grant from Research to Prevent Blindness, New York, NY.
Disclosure: M.L. Gabriele, None; G. Wollstein, Carl Zeiss Meditec (F), Optovue (F), P; H. Ishikawa, P; P.L. Kagemann, None; J. Xu, P; L.S. Folio, None; J.S. Schuman, Carl Zeiss Meditec (F), P