Ординатура / Офтальмология / Английские материалы / The Retina and its Disorders_Besharse, Bok_2011

There is a great deal of information contained in multifocal records that is only partially understood and still largely unexploited. Combining tests of function and structure promises to advance the understanding of both types of data and the pathogenesis of diseases. A case in point is the example of x-linked retinoschisis shown in Figure 6. The availability of OCT data on the different areas of the retina might have greatly helped our understanding of the connection between function and stricture in this and other similar cases.

Acknowledgment

The studies on which this article is based have been supported in part by NIH grant EY06961.

See also: Adaptive Optics; Optical Coherence Tomography; Primary Photoreceptor Degenerations: Retinitis Pigmentosa; Primary Photoreceptor Degenerations: Terminology; Secondary Photoreceptor Degenerations: Age-Related Macular Degeneration.

Further Reading

Hood, D. C. (2000). Assessing retinal function with the multifocal ERG technique. Progress in Retinal and Eye Research 19: 607–646.

Miyake, Y. (2008). Electrodiagnosis of Retinal Diseases. Tokyo: Springer.

Sutter, E. E. (2001). Imaging visual function with the multifocal m-sequence technique. Vision Research 41: 1241–1255.

Sutter, E. E. (1992). A deterministic approach to nonlinear

systems analysis. In: Pinter, R. B. and Nabet, B. (eds.) Nonlinear Vision, pp. 171–220. Cleveland, OH: CRC Press.

Optical Coherence Tomography

W Drexler, Medical University Vienna, Vienna, Austria

ã 2010 Elsevier Ltd. All rights reserved.

Glossary

Adaptive optics (AO) – A technique originally developed for improving imaging performance in astronomy. Applied in vision sciences to correct wave front aberrations introduced by imperfect optics of the human eye to reduce the spot size in the retina and hence transverse resolution of retinal imaging.

Optical coherence tomography (OCT) – Optical analog to ultrasound, for noninvasive threedimensional micrometer resolution visualization of superficial (up to 2 mm) tissue morphology. Optophysiology – Optical analog to electrophysiology that enables noninvasive depth resolved optical probing of retinal physiology.

Optical coherence tomography (OCT) is a rapidly emerging noninvasive, optical diagnostic imaging modality enabling in vivo cross-sectional tomographic visualization of internal tissue microstructure in biological systems at resolution levels of a few micrometers. Novel high-speed detection techniques as well as development of ultrabroad bandwidth and tunable light sources have recently revolutionized imaging performance and clinical feasibility of OCT. In this view, OCTcan now be considered as an optical analog to computed tomography (CT) and magnetic resonance imaging (MRI), not enabling full body imaging, but noninvasive optical biopsy, that is, micrometer/cellular resolution 3D visualization of tissue morphology.

The eye provides easy optical access to the anterior segment and the retina due to its essentially transparent nature. Axial and transverse resolutions are decoupled in OCT. While the axial one is mainly determined by the optical bandwidth of the employed light source, the transverse one is mainly given by the numerical aperture of the optics that is focusing the beam to the tissue. As a consequence, the axial OCT resolution for retinal imaging is not limited by the low-numerical aperture and large depth of focus of the human eye as it does in scanning laser ophthalmosocopy. For this reason, ophthalmic and especially retinal imaging has so far not only been the first, but also the most successful clinical application for OCT. Objectively this is evidenced by the fact that nearly 2500 (49%) of the 5000 OCT publications (including only peer-reviewed articles) have been published in ophthalmic journals. About 1300 (25%) have been published in optical or biomedical journals demonstrating the significant

emphasis on OCT technology development since its invention. In addition, more than half a dozen companies offer this technology in its fourth generation for 3D retinal OCT. Considering the fact that OCT has been introduced only about two decades ago, retinal OCT therefore represents the fastest adopted imaging technology in the history of ophthalmology.

Figure 1 depicts an overview of OCT development with respect to axial resolution and data acquisition speed (measurement time) for morphologic (as opposed to functional) retinal imaging. After its first in vitro demonstration in 1991 and first in vivo imaging studies of the human retina in 1993 (cf. Figure 1(a)–1(d)), OCT has rapidly developed as a noninvasive, optical medical diagnostic imaging modality providing two-dimensional information of retinal structure with resolution about one order of magnitude better than ultrasound. Since its invention, the original idea of OCT was to enable noninvasive optical biopsy, that is, the in situ imaging of tissue microstructure with a resolution approaching that of histology, but without the need for tissue excision and postprocessing. As a consequence, two milestone developments that improved key technological OCT parameters – axial resolution and measurement time – significantly contributed to realize the optical biopsy idea of OCT. A first step toward this goal was the introduction of ultrahigh resolution OCT enabling a noticeably superior visualization of tissue microstructure, for example, all major intraretinal layers as well as cellular resolution OCT imaging in nontransparent tissue by improving axial OCT resolution by one order of magnitude from the 10–15 mm to the (sub)micrometer region, accomplishing 2–3 mm in the living human retina (cf. Figure 1(e)–1(g)). Advances in photonics technology including the development of ultrabroad bandwidth and high-speed tunable light sources as well as high-speed detection techniques have enabled a considerable improvement in data acquisition speed from several hundreds of A-scans/s to 30 000 A-scans/s and recently 300 000 A-scans/s (cf. Figure 1(h)–1(j)).

Despite numerous successful and clinically valuable OCT applications in the anterior eye segment (mainly in the cornea and anterior chamber angle), retinal OCT had a significantly higher clinical impact so far. This is mainly due to the lack of competing noninvasive techniques that can provide comparable wealth of information about the living human retina. Hence, the technological improvements that have been accomplished in the last decade not only enabled unprecedented OCT performance that is unique among noninvasive diagnostic

techniques, but it also established retinal OCT as a state- of-the art noninvasive, complementary ophthalmic diagnostic methodology. For this reason, this article will mainly focus on retinal OCT.

3D Ultrahigh Resolution Retinal OCT

While third generation commercial retinal OCT systems (Stratus OCT) were based on the so-called time domain OCT, enabling up to 400 A-scans/s (one-dimensional (1D) measurements), fourth generation (spectral or Fourierdomain based) commercial retinal OCT systems, nowadays can perform up to 100 times more measurements enabling either highly sampled (high definition) 2D tomograms or 3D (volumetric) imaging of the retina. This is due to an alternative detection method that uses either a CCD camera-based spectrometer (spectral of Fourier domain), that can be read out quickly or a fast tuneable laser (swept source OCT or frequency domain OCT) as compared to a moving mirror (time domain) to perform a single A-scan. The first version is inherently more efficient and enables to raster scan the retina analog to a scanning laser ophthalmoscope (SLO) but thereby not only acquiring tissue information from a single 2D plane (in focus), but the full morphological depth information from an entire volume

with a depth resolution mainly given by the optical bandwidth of the employed light source.

A volumetric data set is acquired for the imaged area that can then be viewed and analyzed in ways similar to those used with CT or MRI scans (cf. Figure 2(a) and 2(b)). Hence, the imaged volume can arbitrarily be cut according to the necessary diagnostic needs, for example, like an ultrahigh resolution SLO in en face (C-mode) tomograms (cf. Figure 2(c) and 2(d)). Figure 2 demonstrates 3D UHR OCT in the foveal region of a patient with retinal pigment epithelium (RPE) atrophy. Note that the axial dimension is twofold enlarged as compared to the other two dimensions for better visualization. The 3D representation of the macular region is presented at different-angled views (cf. Figure 2(a)) depicting the pathological change in the topography of the foveal depression as well as enabling unprecedented views in which the retina can be observed from any direction, including from below (cf. Figure 2(b)). Figure 2(e)–2(h) present virtual biopsy/surgery using 3D UHR OCT in combination with 3D data rendering which allows the user to excise and remove any given layer or part of the retinal volume to visualize intraretinal morphology.

The clinical benefit of these spectral of Fourier domain-based OCT instruments is demonstrated in highly sampled 3D visualization of the retina during a reasonable short data-acquisition time resulting in more

reliable and reproducible 2D thickness maps of (intra) retinal layers. Furthermore, retinal locations are less likely to be missed that are important for diagnosis since an entire volume is measured instead of deciding the location of B-scans during OCT acquisition. This might result in improved diagnosis of retinal pathologies, better understanding of retinal pathogenesis, as well as enhanced (objective) monitoring of novel therapy approaches.

3D Wide-Field Choroidal OCT

So far, commercially available retinal OCT has mainly been performed in the 800-nm wavelength region. This is mainly due to easy availability of broad bandwidth light sources and detectors (CCD cameras) in this wavelength region. Although OCT systems centered at 800 nm can resolve all major intraretinal layers, they only enable limited penetration beyond the retina, due to multiple scattering and absorption in the melanin-rich RPE. This results in limited visualization of the choriocapillaris and choroid. Moreover, in clinical OCT, turbid ocular media (e.g., cataract or corneal haze) represent a significant challenge when imaging the retina. Since scattering in biological tissues decreases monotonically with increasing wavelength, OCT imaging at 1060 nm can deliver deeper tissue penetration enabling delineation of choroidal structure. Being less sensitive to scattering eye media and enabling enhanced penetration into the choroid up to the choroidal–scleral interface at 1060 nm might therefore significantly improve the clinical feasibility of retinal OCT.

2D time domain-based OCT was initiated about 5 years ago demonstrating improved visualization of superficial choroidal structure. With the introduction of more efficient and hence more sensitive spectral of Fourier domain OCT, 3D OCT at 1060 nm demonstrated wide-field visualization of the entire choroid up to the sclera. In this approach, a cost-effective, easy-to-implement system

based on a high-speed InGaAs linear 1024 pixel array (SUI-Goodrich) enabling 47 000 A-scans/s, 5–8 mm axial resolution, and 2.6 mm scanning depth in tissue was developed. Figure 3 depicts the comparison of 3D OCT at 800 nm (cf. Figure 3(a) and 3(b)) versus 1060 nm (cf. Figure 3(c) and 3(d)) in the same normal eye. Enhanced visualization of the choroid up to the choroidal sclera interface is accomplished using 3D 1060-nm OCT as compared to 3D 800-nm OCT (cf. yellow arrows in Figure 3(d)). In the subject with light fundus pigmentation, 3D 1060-nm OCT enables penetration beyond the choroid into the sclera (cf. red arrows in Figure 3(g)). Highspeed 3D 1060-nm OCT also enables en face visualization of the choroidal vasculature without the use of any contrast agent (cf. Figure 3(c), 3(e), and 3(f)).

High-speed 3D 1060-nm OCT therefore now enables unprecedented visualization of all three choroidal layers giving access to the entire choroidal vasculature. In addition, 2D choroidal thickness maps might have significant impact in the early diagnosis of retinal pathologies such as glaucoma, age-related macular degeneration and might contribute to a better understanding of myopigenesis.

Cellular Resolution Retinal OCT

For retinal OCT imaging, the cornea and the lens act as the imaging objective, thereby determining the numerical aperture and hence the beam diameter in the retina. This diameter specifies the transverse OCT resolution that is of the order of 20 mm for a beam of 1-mm diameter (at 800 nm) – approximately one order of magnitude worse than the best axial OCT resolutions accomplished so far. This can be improved by dilating the pupil and increasing the measurement beam diameter. In practice, however, for large pupil diameters, ocular aberrations limit the minimum focused spot size on the retina, even for monochromatic illumination. An alternative and promising

approach is to use adaptive optics (AO), which was originally developed to improve the resolution of astronomical imaging, to minimize ocular aberrations, reduce retinal spot size, and hence to improve transverse OCT resolution.

In ophthalmic AO a wave front sensor measures the individual ocular aberrations of the investigated eye, calculates an inverse wave front, and sends this information to a correcting device – a deformable mirror – for aberration correction. This procedure is performed in real time in a closed-loop configuration. In a couple of tenths of a second, ocular aberrations are compensated and continuously corrected during the entire measurement procedure. In the present approach, a deformable mirror (Mirao52, Imagine Eyes, France) with a unique performance in terms of amplitude ( 50-mm stroke) and linearity was used, allowing for correcting highly aberrated normal or pathologic eyes. Furthermore a 140–160-nm Ti:sapphire laser (Femtolasers Integral, Femtolaser, Vienna, Austria) in combination with a CMOS Basler sprint spL4096-140k camera (Basler AG Germany) enabling 160 000 A-scans/s with 1536 pixels was used, resulting in ultrahigh speed cellular resolution retinal imaging with isotopic resolution of 2–3 mm. Furthermore, the chromatic aberrations of the eye in the 700–900 nm wavelength region of the employed light source have been compensated with a special lens.

The combination of high stroke deformable mirrorbased AO, with ultrahigh speed, ultrahigh resolution OCT employing compensation of the eye’s chromatic aberrations enabled isotropic OCT resolution of 2–3 mm. It is noteworthy that the extremely high measurement speed in combination with sufficient system sensitivity at this speed is essential to maintain cellular resolution morphology information despite motion artifacts. In analogy to AO, scanning laser ophthalmoscopy measurements, AO OCT raster scans a retinal area, but acquires full morphological information as a function of depth in the region of interest, without the need to scan the depth of focal plane. As a consequence, 3D morphology of single photoreceptor (PR) outer segments in addition to cellular microstructure at the level of the RPE and choriocapillaris can be visualized in the living human retina. Figure 4 depicts in vivo cellular resolution retinal imaging in a normal human retina at about 4 parafoveal. A rendered volume of about 350 80 100 mm is presented. En face representations at different depths of the volume reveal cellular resolution intraretinal microstructure.

3D information of intraretinal morphology at cellular level might not only revolutionize ophthalmic diagnosis, but also significantly contribute to a more precise interpretation of OCT tomograms. Despite numerous clinical

ELL

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Figure 4 In vivo cellular resolution retinal adaptive optics OCT. 3D tomogram and en face cross sections of a normal human retina acquired at 4 parafoveal at 160 000 A-scans/s. Various retinal layers in photoreceptor layer may be distinguished with this technique. From top to bottom slice: ellipsoids (ELL), outersegment tips (OS), outer segment tips (OS tips), retinal pigment epithelium (RPE), choriocapillaris (CC), and choriocapillaris/ choroid (CC/CH). Width of slides corresponds to 350 80 mm.

investigations using UHR OCT and systematic studies comparing histology with UHR OCT in in vitro animal models to correctly interpret OCT images, it is noteworthy that a comprehensive, reliable interpretation of all OCT intraretinal layers has not yet been accomplished. Although the state-of-the-art ophthalmic OCT technology enables significantly improved visualization of intraretinal layers, caution is therefore imperative regarding proper interpretation of OCT tomograms, especially of the distal part of the retina. Figure 5 depicts state-of-the-art OCT interpretation of intraretinal layers in a high definition ultrahigh resolution OCT of the optic disk (a) and foveal

(b) region of a normal subject. From the proximal to the distal part of the retina the nerve fiber layer (cf. Figure 5(a), NFL) as well as plexiform layers (cf. Figure 5(a), inner (IPL) and outer plexiform (OPL) layer) appear backreflecting and therefore as a strong signal in the OCT tomogram. The ganglion cell layer (cf. Figure 5(a), GCL) as well as nuclear layer (cf. Figure 5(a), inner (INL) and outer nuclear (ONL) layer) appear less back-reflecting and therefore as a low signal in the OCT tomogram. The external limiting membrane (cf. Figure 5(b), ELM), the inner

and outer PR segments (cf. Figure 5(b), IS PR, OS PR), as well as the choroid (cf. Figure 5(b)) are also properly confirmed by literature. Red-labeled layers indicate that caution is imperative and the correct interpretation needs more conclusive studies to confirm correctness. This applies to the internal limiting membrane (cf. Figure 5(a), ILM) and the distal part of the retina involving probably the distal tips of the PRs – sometimes also referred to Verhoeff ’s membrane (cf. Figure 5(b)), the RPE including Bruch’s membrane (cf. Figure 5(b), RPE/BM) as well as choriocapillaris (cf. Figure 5(b)).

To date, the most distal layer, that is, strongest continuous distal signal in OCT tomograms, has been interpreted as the RPE layer. Although literature describing the light–RPE interaction in the near-infrared region around 800 nm would confirm this interpretation, the relatively thick (up to 30 mm) appearance in OCT tomograms is not supported by histological findings which describe the RPE as a monocellular layer. In addition, the appearance of the RPE and distally adjacent layers as visualized by OCT varies in eyes with different pathologies. Figure 6 depicts the region indicated with a red rectangle in Figure 5 but imaged with AO OCT. 3D information at cellular resolution level significantly helps to correlate OCT findings with well-known anatomy from histology and therefore more precisely interpret 2D OCT tomograms. Hence, from the external limiting membrane (cf. ELM in Figure 6) toward the distal part of the retina, the first bright (white) signal indicates the junction between inner and outer PR segments. En face cellular resolution representation clearly reveals single ellipsoids. After a thicker, less bright (black) layer indicating a part of the PR outer segment, five signal bands with alternating bright (white) and weak (black) signal appearance is revealed by AO OCT. En face information of the first bright (white) band suggests that this layer corresponds to the outer part/outer tips of the outer PR segments. It is noteworthy that this structure as well as the junction between the inner and outer segment of the PR cannot be resolved within a circular zone of less than 2 parafoveal. Their too-tight spacing and the limited numerical aperture of the human eye in addition to weak contrast has not yet allowed the visualization of PRs in the center ( 2 central region) of the fovea. Nevertheless, they are included in this figure for more precise interpretation. The next weak (black) signal band might correspond to the RPE somas, since the adjacent bright (white) signal band is clearly identified to correspond to the RPE layer due to the hexagonal structure of single RPE cells as visualized by AO OCT in the indicated en face view. The next weak (black) signal band might correspond to Bruch’s membrane, although literature describes the thickness of this membrane close to the resolution limits of AO OCT. This interpretation is nevertheless supported by the fact that the cellular resolution en face image of the next bright (white) signal band indicates that it might correspond to

the choriocapillaris. It is noteworthy that despite the superb imaging performance of AO OCT, two layers are still labeled red in Figure 6 indicating that more studies are needed to finally confirm their proper interpretation.

Figure 7 demonstrates the concept of clinical cellular resolution AO OCT in a patient with type 2 macular telangiectasia. A commercial 3D OCT is used to prescreen a larger volume to identify suspicious locations.

These areas are then investigated (at the moment still with a separate system) with AO OCT at cellular resolution level revealing quite normal vasculature and PR appearance at 6 parafoveal whereas there is severe impairment at 0 .

Functional Retinal Imaging Using OCT

Numerous functional OCT extensions have been developed in the past of which Doppler OCT, measuring the blood flow velocity and polarization sensitive OCT, imaging depth resolved tissue birefringence have been the most developed and successfully applied in retinal imaging. Noncontact, depth-resolved optical probing of retinal responses to visual stimulation with 10 mm spatial resolution, achieved using functional ultrahigh resolution OCT, has recently been demonstrated. This method relies on the

observation that physiological changes in dark-adapted retina caused by light stimulation can result in local variations in tissue reflectivity. This functional extension of OCT can be considered as an optical analog to electrophysiology and has therefore been called optophysiology.

To determine the sensitivity of optophysiology for the detection of changes in retinal reflectivity triggered by light stimulation, a dark-adapted living in vitro rabbit retina was exposed to a single flash of white light and optophysiology data were acquired synchronously with electroretinogram (ERG) recordings (cf. Figure 8). Throughout the functional experiments the isolated retinas were stimulated with single, 200-ms long white light flashes (cf. yellow rectangle in Figure 8(c), 8(d), 8(f)–8(h), and 8(j)–8(l)). A morphological B-scan was first taken from the measurement location (cf. Figure 8(a)). Multiple OCT depth reflectivity profiles (A-scans) were then acquired at one transverse location in the retina synchronously with ERG

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recordings. The OCT A-scans were combined to form 2D raw data M-tomograms presenting the retina reflectivity profile as a function of time. The optical data were processed using a cross-correlation algorithm to account for any movement of the retina caused by the solution flow and for calculation of the optical background (average over the pre-stimulation A-scans of each M-tomogram) and generation of differential M-tomograms (cf. Figure 8(b), 8(e), and 8(i)). Optophysiological signals could be extracted from various retinal layers, so that depth-resolved optical back scattering changes that resulted from physiological processes induced by the optical stimulus could be detected. As expected, in the nonstimulated retina (cf. Figure 8(b)–8(d)) the optical reflectivity of the PR layer did not change significantly with time. When the retina is exposed to the light stimulus (marked by the yellow box), changes were seen in optical backscattering at locations corresponding to the inner (cf. Figure 8(f)) and outer (cf. Figure 8(g)) segment of the PR layer which correlated with changes in the corresponding ERG (cf. Figure 8(h)). When potassium chloride (KCL) was applied to the retinal sample to inhibit PR function (cf. Figure 8(i)–8(l)), the optical changes observed in the PR inner segment (IS) and outer segment (OS) of the PR layer were close to the optical background level and showed no correlation to

the onset of the light stimulus. Depolarization of the cell membranes can occur during conduction of an action potential which could be detected by UHR OCT, but also by detection of spatially resolved change in backscattering over time. The exact origin of the detected optophysiologic signals is unclear but might be related to the dipole reorientation (and therefore refractive index changes) at the PR membrane. Alternatively they could arise from lightinduced isomerization of Rhodopsin in the outer PR segment or metabolic changes in the mitochondria of the inner PR segments.

These optophysiological findings in the in vitro rabbit model have recently been tried to transfer to the in vivo human retina. Due to limitations in coordinating stimulus and data acquisition and the very slow time course of PR recovery after photobleaching stimulus, a protocol that included imaging a normal subject at intervals of about 1 min was used. Dark-adapted human eyes were briefly subjected to localized photobleach. For 20 min prior to, and 30 min after stimulus, volumetric optical coherence tomograms were collected partially overlapping with the bleached region. A location at the peak rod density at 11 temporal was scanned because rod responses are interesting in regards to dark adaptation and rod responses had been successfully observed the rabbit in vitro experiments.

The scanned patch was sampled at 512 A-scans in the long axis of the rectangle (fast axis) and 256 samples in the narrow axis (slow axis). Given the 47-kHz A-scan rate of the camera, this volume could be collected in 2.8 s. Twenty double volumes were completed before the subject was exposed to a white, 48 000 cd m–2 stimulus, bleaching over 99% of both photopic (duration 6 s) and scoptopic photopigments. As the eye recovered dark accommodation, 35 more double volumes were recorded at 1 min intervals. Tomograms were segmented into retinal layers by a newly described algorithm exploiting information in adjacent B-scans. En face fundus images extracted from major

intraretinal layers were laterally registered manually. Time series summarizing the observed backscatter in selected layers for the bleached and unbleached areas are shown with a variety of corrections and normalizations applied: tomograms were corrected for inherent sensitivity roll off; the ratio between other layers and an assumed unchanging layer (RPE) as well as the ratio of stimulated area to unstimulated area were calculated.

Figure 9 depicts results obtained from one normal subject. For each tissue layer in each measured volume, the result of the experiment is summarized by the mean voxel intensity underneath the retinal surface receiving the

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