2.1.

Optical Design

The AOSLO system has a double-pass optical path, including a light source, horizontal and vertical scanners, a wavefront sensor, a DM, and a photomultiplier tube (PMT) detector. Note that the wavefront sensor and PMT detector are on-axis optical modules that are relatively free from aberrations. Only the light delivery path is optically designed and optimized using ZEMAX (Radiant ZEMAX LLC, Redmond, Washington), and an overview of the optimized optical design is shown in Fig. 1. Spherical mirrors in an off-axis arrangement are used to relay the planes conjugate to the pupil of the eye to avoid back-reflected light, and these mirrors are used in pairs, creating an afocal telescope to preserve conjugate planes. To form a rectangular imaging raster on the retina, a resonant scanner (SC30-3 $\times 4$, Electro-Optical Products Corp) line-scanned the retina horizontally at 16 kHz with an optical aperture of 3 mm, and a slow galvometric scanner (6800HP, Cambridge Tech. Corp) scanned the vertical direction at 30 Hz.

Fig. 1

(a) A top view schematic of the AOSLO optical design. Note that optical components are folded in three dimensions and is flattened for illustration. All optical components are labeled and described in the legend. The optical design is the light delivery path excluding the on-axis optical components free from aberrations. (b) The light delivery path in side view as shown in Fig. 1(a). M5 and M6 are arranged out of the main optical plane to compensate for astigmatism.

A 670-nm superluminescent diode (SLD) with 10-nm bandwidth (SLD-261-HP, Superlum, Ireland) was used for wavefront sensing and retina imaging after being collimated to produce parallel lights with a diameter of $\sim 6.0\mathrm{mm}$ through the collimating lens L1. The delivery optical path consists of four $4\text{-}f$, afocal telescope pairs that relay the eye pupil to several pupil planes, including the planes of the DM, horizontal and vertical scanners, and wavefront sensor. The DM is placed at the pupil conjugate closest to the light source to avoid ray light movements or wanders in the DM surface. All the afocal telescopes of spherical mirrors are ordered to achieve matches between each pupil aperture. Since the off-axis mirrors are likely to generate off-axis aberrations such as astigmatism and coma, a straightforward way to reduce the off-axis angle and astigmatism is to use spherical mirrors with long focal lengths, but this increases the overall size of the device. By folding spherical mirrors in a nonplanar configuration and alternating the angle of beam reflection of these mirrors from the horizontal and vertical directions, off-axis aberrations can be minimized.²⁸ Therefore, spherical mirrors M5 and M6 were carefully used off the optical plane to compensate astigmatism aberrations, and folding angles were optimized (under constraints for ray clearance and mechanical mount clearance) on both the horizontal and vertical directions to minimize monochromatic aberrations in the retina.

In addition, the spherical mirror M8 prior to the eye consisted of 25.4-mm-diameter lenses with an effective focal length of 160 mm. This focal length was chosen to give a working distance of $\sim 30\mathrm{mm}$ and have sufficient room for the plano mirror PM2 to fold the light path for easy retina imaging. Limited by constrains of XY scanners, a $3\mathrm{deg}\times 3\mathrm{deg}$ FOV was used to design the AOSLO. However, the real FOV for retina imaging should be cropped smaller than 2 deg, suffering from the human eye’s isoplanatic angle in an AO system.

After the initial optical structure was chosen, the optical design was optimized by minimizing spot sizes on the retina for 17 uniformly distributed points spanning the FOV and for three different vergences ($-5\mathrm{D}$, 0D, and $+5\mathrm{D}$). Taking a modified Gullstrand eye model,³³ diffraction-limited performance for a vergence range of 10D over all the FOVs has been achieved for high-resolution AO imaging, and the optical optimization performance is shown in Fig. 2. Note that all spot diagrams were calculated on the focal plane of the model human eye, assuming an AO correction of low-order aberrations when a seventh-order Zernike polynomial was adopted to estimate the wavefront aberrations in ZEMAX. For an emmetropia eye, much better performance can be expected that the residual wavefront root mean square (RMS) was lower than $\lambda /21$. With the increase of vergence, the flatness of AO corrections would decrease for myopic and hyperopic eyes with low-order aberrations, and the residual wavefront RMS of astigmatism would slightly increase while still meeting the diffraction limit requirement. Benefitting from the large stroke of the bimorph DM to accomplish AO aberration corrections, the residual wavefront RMS at 670 nm was lower than $\lambda /14$ for all vergences.

Fig. 2

Spot diagrams for nine typical configurations evaluated on the retina over a $3\mathrm{deg}\times 3\mathrm{deg}$ FOV for a vergence range of 10D in the AOSLO optical design. Configurations are grouped by vergences, and all configurations are diffraction limited for the wavelength of 670 nm with a pupil diameter of 6.4 mm. The radius of Airy disk (black circle) is $2.09\mu \mathrm{m}$.

Since the horizontal and vertical scanners move to form the two-dimensional (2-D) imaging raster at the retina plane, measures must be taken to ensure that the chief ray pivots around a stationary point in the center of all the pupil planes in this AOSLO system, including the two scanners and the eye pupil. Figure 3 shows light maps on the three pupil planes. Along the light passing to the retina, the aberration is growing to cause the displacement of the chief ray on those pupil planes, and the maximum beam wander for all the FOVs was $\sim 6.25\%$ and occurred on the eye pupil. Actually, beam wander in the pupil planes is bound to occur due to the mirror-based reflective optical system with beam scanning, but this is acceptable considering that human pupils can usually be dilated to 7 to 8 mm, and the beam wander will decrease once the scanning FOV is appropriately reduced.

Fig. 3

Light maps for the three pupil planes over all the FOVs, and different scanning FOVs are coded by color. Black circle represents the maximum diameter occupying all the wandered rays.

For those on-axis optical components, a custom-made Hartmann–Shack wavefront sensor (HSWS) consisting of a square lenslet array with a focal length of 5 and $0.2\mathrm{mm}\times 0.2\mathrm{mm}$ pitches was used for wavefront sensing. To complete the double-pass optical path, light backscattered from the retina passed through the delivery path and was split by several beam splitters (not shown in this figure), of which an 80/20 beam splitter was utilized to separate the illuminating and imaging beam. Then a 95/5-beam splitter was ordered to separate the detecting and wavefront sensing beams. The detecting beam was collected using PMTs (H7422-20, Hamamatsu, Japan) and then focused by an achromatic lens with a focal length of 200 mm. For the confocal configuration, a pinhole with a diameter of $50\mu \mathrm{m}$ equal to $\sim 1.44$ times the Airy disk diameter was set in front of the PMT, which has been demonstrated to give a good balance between image sharpness and throughput.^34,35

2.2.

Adaptive Optics Control and Electronic Design

Compared to other widely used clinical SLO devices, AO for aberration correction was integrated with the SLO module in an increase of complexity manner, thereby making it difficult to operate on separate computers allowing for aberration correction, scanning timing, and retina imaging.³⁶ Therefore, we optimized the AO control and electronic design to realize an easy-to-use AOSLO system for aberration correction and retina imaging.

In a typical implementation of a wavefront sensor-based AO system for human retina imaging, the limited stroke of DM made it difficult to accurately compensate for ocular aberrations, and additional trial lenses were introduced to precompensate low-order aberrations with large amplitudes.^37,38 The DM played the role of high-order aberration correction, which in general requires lower stroke. However, this approach requires manually changing the graduation of trial lenses for each individual subject, which can be a tedious and complex process^28,39 and can even disturb the AO aberration correction once the trial lens is located away from the pupil conjugate. In this study, with the developed bimorph DM of large stroke values (up to $50\mu \mathrm{m}$), we propose a successive AO control strategy to compensate ocular aberrations without additional trial lenses, which is a bit similar to the strategy of the woofer–tweeter dual DM AOSLO.⁴⁰ However, only the single bimorph DM, rather than two DMs, is utilized for lower- and higher-order aberration corrections. In practice, the ocular aberrations are measured through the HSWS, and efficiently partitioned to the two parts of low-order and high-order aberrations in the process of wavefront restorations. Therefore, the bimorph DM is predistorted in an open-loop mode to match the expected wavefront of low-order aberration correction, in which all actuators subjected to the same voltage will produce a defocus phase, and different voltages on the fan-shaped actuators can produce an astigmatism phase. Once the open-loop AO correction of low-order aberrations works well after two to five iterations ($\sim 20$ to 50 ms), the closed-loop AO control is engaged to correct the high-order aberrations.

In the electronic module of the AOSLO system, the HSWS images were captured with a speed of 120 Hz using a GigE camera (Manta G-235, Allied Vision, Germany), and then AO control signals were calculated and updated in sequence, including wavefront errors, aberration deflection maps, Zernike coefficients, and the DM compensation signals. When the desired focus of the retina is achieved by open-loop correction of low-order aberrations, such as near the plane of the photoreceptors, the closed-loop AO control is used to minimize the wavefront errors.

When the AO aberration correction works well and is locked, an independent real-time controller (custom-built field-programmable gate array) generates driving signals and synchronous position signals for the horizontal and vertical scanners, and start-of-frame synchronization signals are also provided to the frame grabber (PCIe-1433, NI) to capture retinal high-resolution images from the PMT detector. Unlike other AO retinal imaging systems that require separate computers to control different subsystems,^36,38 only one computer is needed to implement the two electronic control modules in the proposed AOSLO system, which is shown in Fig. 4. Retinal images in AVI image files with a speed of 30 frames per second can be recorded and stored in the database of the computer, and image processing, such as montaging, is performed offline in MATLAB (MathWorks, Inc.).

Fig. 4

The electronic control of the AOSLO system. Dashed lines represent the two major control subsystems, the high-resolution imaging (left), and the AO for aberration correction (right). Red lines with arrows present the optical signals, and the black lines present the electronic signals.

2.3.

Mechanical Design

To reduce the complexity of constructing an AOSLO system, the mechanical design for the system was developed in SolidWorks (Dassault Systèmes SolidWorks Corp., Concord, Massachusetts), as shown in Fig. 5(a). A brief rule is that all the optomechanical components are set on a thick bottom plate, and all the other components such as electronic devices are included in the lower part of the system for weighting balance. According to the maximum permissible positional errors of optical design, a mechanical tolerance analysis was performed to determine the adjustments of each mirror and lens to optimize the assembly, and then all the mirror mounts, lens tubes, and lens spacers were designed and fabricated to accommodate the closely spaced optics of the system and minimize its footprint. Except for the bottom plate made of thickened iron, the internal skeleton and other structural components were made of aluminum to simplify fabrication and maintain low weight. To optimize performance between the assembly and operation, we used the minimum number of adjustments in the mechanical design, including the adjustability of the confocal pinhole, horizontal and vertical scanners, HSWS, and DM. The confocal pinhole was mounted on a linear motion stage with total travel lengths of 25 mm to allow adjustments in the $XYZ$ direction. The horizontal and vertical scanners were mounted via a mechanical base with curved slots to adjust the angular orientation while maintaining the correct center of rotation. The HSWS and DM were constructed with three pull/push structures, allowing the tilt/pitch of the laser beam to be adjusted by 3 mm and 15 deg. The outer casing was made of iron with painting and chamfering, and the fabricated setup is shown in Fig. 5(b). A chin rest was mounted into the bottom plate using $XYZ$ translation stages for light beam alignment. The AOSLO prototype weighed 55 kg and was 80-cm long × 60-cm wide × 120-cm tall. However, the AOLSO mainframe had a much smaller size of 80-cm long × 60-cm wide × 20-cm tall and a lighter weight of 18.5 kg when all other electronic devices were moved to an additional box.

Fig. 5

The mechanical design of the AOSLO system in top view (a) showing the internal optical components. (b) Photograph of the fabricated setup with case.

3.1.

Adaptive Optics Performance

Prior to retina imaging, all human subjects were fully informed of the procedures and consequences of the study, and all subjects provided signed informed consent. Eyes to be imaged were dilated with 0.5% tropicamide to 6- to 7-mm pupils and illuminated by the SLD laser beam always kept below $140\mu \mathrm{W}$, matching safety standards of the American National Standards Institute maximum laser exposure in more than one order of magnitude, even if the exposure time lasts almost 50 min.^41,42 In practice, each single imaging experiment can be only executed for a period of no more than 3 min or even shorter time, due to eye motion or blink.

Subjects had refractive errors up to 6.5D, with lower-order aberrations of defocus and astigmatism. Such large, lower-order aberrations can lead to the failure of closed-loop control, especially since the large defocus aberrations take many cycles to be corrected while causing the correction to oscillate. To prevent the AO aberration correction from failing, the bimorph DM was initiated first and run in open-loop mode, where it corrected the system and subject defocus aberrations through a fixed and static aberration correction until the wavefront aberration RMS was reduced to below $1.0\mu \mathrm{m}$, ensuring that the HSWS sensor was within its working range at the start of the closed-loop correction for high-order aberrations. Then the HSWS was activated to measure high-order aberrations, and the DM was left in closed-loop mode for a dynamic AO correction until the wavefront RMS was decreased to a diffraction-limit performance.

The estimation of the wavefront aberrations was fitted to a seventh-order Zernike polynomial, using the normalization given by the OSA standard for ocular aberrations.^43,44 Figure 6 shows an example of the AO performance achieved in one healthy human subject (male, age 32, pupil with a diameter of $\sim 6.42\mathrm{mm}$). In Figs. 6(a)–6(c), color-coded aberrations maps are illustrated for three cases: initial aberrations without AO correction, aberrations with open-loop correction, and aberrations with closed-loop correction. The averaged aberrations separated by Zernike order and wavefront RMS reduction course with the AO correction are also shown in Figs. 6(d) and 6(e). Note that the initial aberrations presented typical myopia ($\sim 2\mathrm{D}$) and astigmatism ($\sim 1\mathrm{D}$) and a rather irregular cornea that introduced some high-order aberrations. With open-loop AO correction, the average RMS error improved from 2.16 to $0.77\mu \mathrm{m}$, mostly reducing low-order aberrations. After the closed-loop AO correction, the RMS errors generally reduced to $0.034\mu \mathrm{m}$. The improvement in all Zernike orders is shown in Fig. 6(d), although the most improvement was seen in the first six orders. In Fig. 6(e), the 10% to 90% rise-time for the AO correction was $\sim 100\mathrm{ms}$, which corresponds to a temporal bandwidth of $\sim 10\mathrm{Hz}$ for the AO correction duration.

Fig. 6

Example of the AO performance for one healthy subject with a 6.42-mm pupil. Wavefront aberration maps in (a)–(c) represent the initial aberrations, aberrations with open-loop correction, and aberrations with closed-loop correction, respectively. The bar diagram in (d) shows the aberration coefficients in Zernike orders, and the temporal profile in (e) shows the RMS wavefront error reduction during the whole AO correction.

The AO performance was further studied in a small-scale population experiment. A total of 20 subjects between the ages of 20 to 40 years (average: 32 years; standard deviation: 6 years) were recruited for the experiment. Figure 7 shows the RMS wavefront errors with and without AO corrections. All the subjects were corrected to better than $0.08\mu \mathrm{m}$ RMS errors. The average RMS error improved from $3.05\pm 1.17$ to $0.067\pm 0.012\mu \mathrm{m}$ with AO correction.

Fig. 7

Aberration RMS errors of 20 subjects without and with AO corrections in black color bar and blue color bar, respectively. Low-order aberrations are illustrated in red color bar.

3.2.

Photoreceptor Imaging

Since the bimorph DM-based AO was effective in correcting human ocular aberrations, Figs. 8(a) and 8(b) compare retina images without and with AO corrections for a female healthy subject, aged 26 years with a $\mathrm{\Phi }6.84\text{-}\mathrm{mm}$ pupil. Both single-frame images contained $512\text{pixels}\times 449\text{pixels}$ of the same retina location 1.5 deg from the fovea center. With an open-loop correction of relatively large low-order aberrations ($-4\mathrm{D}$ defocus and $-1.5\mathrm{D}/065$ astigmatism), even the best-focus image in Fig. 8(a) cannot distinguish any retinal photoreceptors. When the closed-loop AO correction was activated to dynamically correct ocular aberrations, an image of the same subject and retinal location was shown in Fig. 8(b), and photoreceptor cones can be clearly resolved due to the increased brightness, enhanced contrast, and improved resolution. With AO correction in Fig. 8(b), the ocular wavefront error was corrected to the RMS value of about $0.046\mu \mathrm{m}$, and much sharper structural details of the retina, such as photoreceptors, can be clearly distinguished.

Fig. 8

Comparison of retina images (a) without and (b) with AO closed-loop corrections. (c) The spectra power plot for retina images. The purple dot-dash lines were added to show the trends of the power spectra curves. All single-frame images are $512\text{pixels}\times 449\text{pixels}$, and the retina image without AO has been scaled in intensity $2\times$ relative to the image with AO correction. Both images are taken from the same retina location 1.5 deg from the fovea center, and the FOV subtends $1.7\mathrm{deg}\times 1.5\mathrm{deg}$. The scale bar is $50\mu \mathrm{m}$. Movie sequence of 30 fps is also presented to show the AO correction procedure (Video 1, 2.75 MB, mp4 [URL: https://doi.org/10.1117/1.NPh.6.4.041111.1]).

The power spectra of retina images are presented to demonstrate the improvement of AO correction in Fig. 8(c), where the Fourier spectrum for each frequency is radially integrated. Benefitting from the AO correction of human ocular aberrations, the power spectra increased at low frequencies and decreased at high frequencies. This indicates that the image resolution corresponding to low frequencies was improved, and the noises corresponding to high frequencies were suppressed.

For comparison in this power spectra, the highest resolvable spatial frequency of the retina image was about 75 cycles/deg without AO correction, and it increased to 140 cycles/deg with AO correction. The spatial frequency around the Yellot’s ring⁴⁵ determines the highest resolvable spatial frequency. Highlighted by an arrow in Fig. 8(c), the radius of the Yellot’s ring is about 140 cycles/deg, which is relatively close to the highest resolvable frequency with AO correction. According to the Raleigh criteria, the minimum resolution for a $\mathrm{\Phi }6.84\text{-}\mathrm{mm}$ pupil at the wavelength of 670 nm is about $2.03\mu \mathrm{m}$ (145 cycles/deg in spatial frequency) for a modified Gullstrand eye model,³³ so the retina image with AO correction is very close to but still below the Raleigh diffraction limit, suffering from the system noise of intensity conditions.

3.3.

Axial Sectioning and Imaging

With the bimorph DM of large strokes of $\sim 50\mu \mathrm{m}$, ocular aberrations can be corrected to achieve the diffraction-limit resolution, and enough strokes remain to allow for axial sectioning of the retina. When the AO correction is accomplished to achieve a wavefront error RMS of $<0.1\mu \mathrm{m}$ (can be thought as a successful aberration compensation in most situations), an additional bias to the defocus coefficient is applied to the bimorph DM, and light beam will be focused at different retina depth. Benefited from the confocal pinhole, only light from near the plane of best focus gets imaged, and true axial sectioning of the retina can be obtained. Figure 9 shows retinal images acquired at several depths while continuously changing the DM sectioning depth, which have the same location in the subject of Fig. 8. The focus was initially at the photoreceptor layer in Fig. 9(a), and then sectioned to the blood vessel layer in Fig. 9(c) and the nerve fiber layer in Fig. 9(d). Therefore, three layers of the retina section can be clearly revealed in Fig. 9, which entirely spans $\sim 300\mu \mathrm{m}$ of the axial depth in this retina. For practical purposes, retinal image between the photoreceptor layer and the blood vessel layer is also presented in Fig. 9, where the photoreceptor mosaic cells and the blood vessel are recognizable in Fig. 9(b).

Fig. 9

Axial sectioning and imaging at different retinal depths for the same location as in the subject of Fig. 8. (a)–(d) The light beam is focused at different surface of the retina while continuously changing the DM sectioning depth. (a) Layers of photoreceptor, (c) blood vessel, and (d) nerve fiber layer are clearly distinguished, and retinal image between the photoreceptor layer and the blood vessel layer is also presented in (b). Scale bar is $50\mu \mathrm{m}$. Movie sequence of 30 fps is also presented to show the axial sectioning procedure (Video 2, 13.9 MB, mp4 [URL: https://doi.org/10.1117/1.NPh.6.4.041111.2]).

3.4.

Retinitis Pigmentosa Disease Imaging

The proposed AOSLO prototype was then used to image retinal patients with RP. RP is a genetic disorder of the eyes in which pigments gradually form in the retina, thereby causing the progressive loss of rod photoreceptor cells and further loss of cone photoreceptor cells, eventually resulting in severe vision loss.^46,47 Unfortunately, the loss of photoreceptor cells cannot be imaged by commercially available retinal ophthalmoscope devices due to lack of resolution, and even the most extensive examinations for RP diagnosis, including electroretinograms, visual field testing, and genetic testing, cannot image photoreceptor loss directly.⁴⁷ With the help of the latest commercial AO device, the AO-fundus camera (RTX1, Orsay, France), clinicians can diagnose RP for the first time by observing the absence of photoreceptor cells. Figure 10 shows a representative AO image of a patient with RP (aged 20 years, female). The high-resolution AO-fundus image in Fig. 10(a) shows an abnormality on the photoreceptor morphology [oculus dexter (OD), recorded at location: temporal = 4.0 deg; superior = 3.0 deg], and the AOSLO images of this subject at the same location are presented in Figs. 10(d) and 10(e).

Fig. 10

High-resolution AO images of a female patient with RP at the age of 20. (a) Commercial AO-fundus image of photoreceptor cells with a FOV of $6\mathrm{deg}\times 4\mathrm{deg}$ (OD, temporal = 4 deg; superior = 3 deg). (b), (c) Enlarged regions of interest from the AO-fundus image indicated with the red squares in (a) (overlaid in red are the photoreceptor loss that were manually segmented by several experienced doctors). (d), (e) Single original images ($1.7\mathrm{deg}\times 1.5\mathrm{deg}$) acquired with the AOSLO prototype at the same locations of (b) and (c), respectively (overlaid in red is the photoreceptor loss that was manually segmented by several experienced doctors). Scale bar is $50\mu \mathrm{m}$.

High-resolution images in Fig. 10(a) appear sharp with AO correction, and an abnormality on the photoreceptor morphology can be observed, in which the photoreceptor loss can be noted compared to the normal retina. The enlarged regions of interest in Figs. 10(b) and 10(c) indicated with the red squares in Fig. 10(a) show areas of photoreceptor loss at those locations (cf. overlaid manual segmentation). However, photoreceptor mosaic is slightly blurred, and the vessel edge is particularly blurred in Figs. 10(a)–10(c). The AOSLO images in Figs. 10(d) and 10(e) avoid photoreceptor blurriness, clearly showing the area of photoreceptor mosaic loss (red-labeled areas). Actually, photoreceptor loss areas in Figs. 10(d) and 10(e) are smaller and less than those in Figs. 10(b) and 10(c), and more accurate photoreceptor density and spacing can be concluded in AOSLO images. In Fig. 10(d), the photoreceptor blurriness affected the shape of vessels, causing obscure vascular boundaries. Much more photoreceptor blurriness can be observed in Figs. 10(b) and 10(c), and some are overlaid on vessels (yellow-labeled areas), probably leading to incorrect observations of blood vessels. Correspondingly, there is almost no photoreceptor blurriness in Figs. 10(d) and 10(e), and an apparent bifurcation of vessels can be distinguished clearly, which is mistaken as visible photoreceptor loss in Fig. 10(c).

The Fourier spectral power plots of this subject are shown in Fig. 11. For the AO-fundus images of this subject in Figs. 10(b) and 10(c), the highest resolvable frequencies were about 90 and 95 cycles/deg, respectively. According to the Raleigh criteria, the theoretical minimum resolvable spatial frequency for a $\mathrm{\Phi }\text{-}6.62\text{-}\mathrm{mm}$ pupil at the wavelength of 850 nm is 112 cycles/deg ($2.66\mu \mathrm{m}$ diffraction-limit resolution), so the AO-fundus images in Figs. 10(b) and 10(c) were relatively close to the Raleigh diffraction limit, and the photoreceptor blurriness of AO-fundus images is not due to wavefront errors in the AO correction but rather due to the lack of resolution.

Fig. 11

The spectral power plot for retina images in Fig. 10. The green dot-dash lines were added to show the trends of the spectral power curves.

For this subject at the same location, the highest resolvable spatial frequencies in Figs. 10(d) and 10(e) were about 140 and 145 cycles/deg, which were close to or even better than the Raleigh diffraction limit of 141 cycles/deg ($2.10\mu \mathrm{m}$ diffraction-limit resolution for the 6.62-mm pupil size). It is obvious that the resolution of the AOSLO images is much higher than AO-fundus images under the same conditions. The AOSLO utilized a shorter laser wavelength for illumination than the commercial AO fundus, and the bimorph DM-based AO achieved an aberration compensation close to the diffraction limit. Owing to the improvement of resolution in this AOSLO prototype, many more details in retina diseases will be resolved.

Early signs of RP include loss of photoreceptors and changes in retinal vessels, especially thinning of small vessels and capillaries that are likely to develop in the inner retinal layers. Actually, conventional flood-illuminated imaging does not allow for axial resolution, even if AO is used in the AO-fundus camera RTX1 to correct aberrations, but the contrast is still reduced by light scattered from other retinal layers. Therefore, the AOSLO has the confocal pinhole for axial sectioning, which allows only light from near the plane of best focus to get imaged. Corresponding images of three retinal layers of this RP patient are shown in Fig. 12, and the sections reveal different retinal surfaces 1.8-deg superior to the fovea. Typical images of photoreceptor in Fig. 12(a) indicated that photoreceptors or pigments are concentrated in the front of the blood vessels, covering part of the blood vessels, or along the blood vessels, and most of them are found in the branches of the blood vessels (red-labeled ellipse areas). In contrast to an image where the focus is set to the photoreceptor layer, most vessels appear bright in Figs. 12(c) and 12(d) with the focus set to the superficial layers, because light from the highly scattering blood constituents and light reflection from the top vessel wall can now be collected by the confocal pinhole for imaging. In general, very small capillaries [size similar to the cone photoreceptors, which have been indicated by the red arrows in Figs. 12(c) and 12(d)] can be resolved with the axial section, and these characteristics can be observed to provide more information for retina pathology. In addition, it should be noted that RP causes a narrowing of retinal vessels (indicated by red lines) in inner layers of Fig. 12(c).

Fig. 12

Retinal images acquired at several retinal depths of the RP patient in Fig. 10. (a) Photoreceptor layer (overlaid in red-labeled ellipse are the photoreceptors or pigments concentrations that were manually segmented by several experienced doctors). (b) Between photoreceptor and blood vessel layers. (c) Blood vessel layer (overlaid in red arrows are small capillaries with size similar to the cone photoreceptors, and red lines are added to indicate a narrowing of retinal vessels cause by RP). (d) Nerve fiber layer (overlaid in red arrows are small capillaries with size similar to the cone photoreceptors). The scale bar is 50 μm.

High-resolution and high-contrast imaging of several retinal layers can be achieved using this AOSLO, and small vessels can be resolved to provide more information for pathology. In practice, the added refractive power for axial sectioning were $-0.1\mu \mathrm{m}$ (0.09D), $-0.2\mu \mathrm{m}$ (0.25D), $-0.5\mu \mathrm{m}$ (0.58D), and $-0.8\mu \mathrm{m}$ (0.9D), respectively, for these images in Fig. 12. The estimated thickness between the photoreceptor layer and the nerve fiber layer was $\sim 215\mu \mathrm{m}$. In general, the axial sections of RP appear a little dark and show typically weak contrast (retinal images in Fig. 12 have been promoted in intensity), which is due to the weak scattering of light for AO imaging and can be greatly increased through an offset pinhole configuration⁴⁸ or through a fluorescein angiography AOSLO setup in the future work.