We present an alternative approach for an adaptive optics scanning laser ophthalmoscope (AO-SLO). In contrast to other commonly used AO-SLO instruments, the imaging optics consist of lenses. Images of the fovea region of 5 healthy volunteers are recorded. The system is capable to resolve human foveal cones in 3 out of 5 healthy volunteers. Additionally, we investigated the capability of the system to support larger scanning angles (up to 5°) on the retina. Finally, in order to demonstrate the performance of the instrument images of rod photoreceptors are presented.
© 2012 OSA
Adaptive optics scanning laser ophthalmoscopy (AO-SLO) is a technology which provides high resolution (diffraction limited) images of the human retina [1–3]. AO-SLO is capable to resolve individual cells (e.g. cone photoreceptors, leukocytes) within the human retina in vivo [1,4–7] and a variety of different diseases have been investigated [8–14]. In some diseases even retinal pigment epithelium (RPE) cells could be visualized . Recent overviews of this technique can be found in Refs. [2,16,17]. It should be noted, however, that individual cells (e.g. photoreceptors at a certain eccentricity from the fovea) can be resolved even without the use of adaptive optics . This can be achieved because the light backscattered from the photoreceptors is directional  and the spacing between photoreceptors increases with eccentricity from the fovea . Additionally, the aberrations within these systems are kept low and the influence of aberrations introduced by the optics of the eye is reduced using either trial lenses or small (2-4mm) imaging beam diameter [19,21,22]. Nevertheless, in order to resolve foveal cones or rod photoreceptors the use of AO is required.
AO corrects aberrations that are introduced to the imaging beam by imperfections of the eye optics. In the most commonly used configuration of AO, the wavefront exiting the eye is measured by a Shack-Hartman wavefront sensor (SHWS) and is corrected using one or more correcting devices (e.g. deformable mirrors) that are placed in the imaging path. Hereby, it is essential that the pupil plane of the eye is imaged onto the correcting device and the SHWS, which is normally achieved by using telescopes based on spherical mirrors. (This configuration is also commonly used in AO optical coherence tomography (AO-OCT) [23–27].)
However, it has been shown that aberrations introduced by the imaging system itself may degrade resolution and therefore prevent true diffraction limited imaging . By arranging the spherical mirrors in a very special way, these aberrations can be minimized [29,30] leading to outstanding images of foveal cones  as well as of rods .
Although several parameters are used to characterize the resolution of an AO-SLO system (e.g. theoretical resolution, residual wavefront error) the achieved resolution on the retina is rather difficult to determine. Even though the measured residual wavefront error indicates a diffraction limited performance, the resolution might be lower because the wavefront sensor will not sense all aberrations (e.g. in the case of scanning). The size of rods and foveal cones is just above diffraction limit. Therefore, probably the ability of a system to resolve these structures can be regarded as an indicator for diffraction limited performance.
In this paper we present an alternative approach for AO-SLO. Instead of using spherical mirrors the imaging optics of the system is based on lenses. The aberrations introduced by such a system should be very low. Additionally, a lens based system can in principle be built very compact. First we compare the performance using Zemax simulation of afocal telescopes (AT) based on spherical mirrors with AT based on lenses. Then we demonstrate the performance of the system by imaging the cone mosaic of healthy volunteers in the fovea region. In addition we investigate the imaging performance of the new system in the case of larger area scans (up to 5 degrees) on the retina. Finally, we present images of the rod mosaic recorded with the new instrument.
2. Lens based adaptive optics
In an operational adaptive optics system, the wavefront at the pupil plane of the eye has to be precisely imaged onto a wavefront sensor (SHS), a correcting device (e.g. deformable mirror) and, in the case of a scanning instrument (e.g. SLO or OCT), onto the scanners. Most AO scanning instruments use spherical mirrors for this purpose, mainly to avoid backreflections from lens surfaces which would disturb the wavefront measurement. In order to investigate the differing performance between lens-based telescopes and reflective telescopes we simulated the optical paths using the optical design program ZEMAX (Optima Research, UK) at a design wavelength of 840nm. Let us first assume a simple afocal telescope consisting of two spherical mirrors that are separated by the sum of their focal lengths. The radius of curvature for both mirrors shall be 400mm and the incidence angle (in respect to the optical axis) onto the mirrors shall be 5 degrees (in a planar configuration). The main aberration introduced by this configuration will be astigmatism. Using an entrance pupil of 7mm and a scanning angle of ± 1 degrees we obtain the corresponding spot diagrams in the imaging plane of the telescope as shown in Fig. 1(a) . The black circles indicate the Airy disc and diffraction limited performance. Although the scanning angle is rather small, some of the spots lie outside the Airy disc indicating close to diffraction limited performance. The situation can be improved, as demonstrated previously [29–31], by using a non-planar folding of the telescopes (c.f. Figure 1(b)). However, similar improvement can be obtained if lenses are used instead of spherical mirrors. Figure 1(c) shows spot diagrams in the imaging plane of a lens based telescope consisting of two achromatic lenses with a focal length of 200mm and a diameter of 25mm. (The lenses have been aligned in order to fulfill the condition of a separation between the lenses of 2f). Comparable performance as the folded configuration using spherical mirrors can be observed. The mean RMS of the wavefront of all scanning angles measured at the image plane of the three telescopes was calculated with 0.096µm, 0.023µm and 0.018µm (for the planar, non-planar and lens based configuration), respectively. The folded and the lens based configuration yield a factor of 4-5 lower residual wavefront RMS compared to the normal plane configuration.
3. Experimental setup
A scheme of the experimental setup is shown in Fig. 2 . The light from a superluminescent diode (Superlum, Russia) with a center wavelength of 840nm and a bandwidth of 50nm is collimated (beam diameter of 4.0 mm) and traverses a Glan-Thompson polarizer and a polarizing beam-splitter. (The polarizer is needed because the polarizing beam-splitter provides only an extinction ratio between the two linear polarization states of 1:100, the polarizer provides an extinction ratio of 1:100000). The linear polarized light traverses the first telescope (two achromatic lenses, L1 and L2 in Fig. 2, each with 200mm focal length) and is reflected at the resonant scanner (Cambridge Technologies, Lexington, MA, USA, 8kHz resonant frequency). The second telescope (L3 and L4, again two achromatic lenses with 200mm focal length) images the pivot point onto the galvanometer scanner (Cambridge Technologies, Lexington, MA, USA). The third telescope (L5 an L6, 75mm and 250mm focal lengths, respectively) increases the beam diameter and images the pivot point of both scanners onto the deformable mirror (Mirao 52, Imagine Eyes, Orsay, France). The forth telescope (L7 and L8) decreases the beam diameter down to ~8 mm and images the pivot point and the DM plane onto the pupil plane of the eye. Before entering the eye, the light beam traverses a pellicle (where the fixation light is coupled into the instrument) and a quarter wave plate (QWP, oriented at 45° to the input linear polarization plane). Therefore the eye is illuminated with circular polarized light. The power at the cornea of the eye is kept below 700µW which is much lower than the permissible exposure limits given in the European Laser Safety Standards for a scanning beam (1°x1° scanning angle). The light is attenuated when passing through the system (starting from the polarizer to the exit of the system) by a factor of ~2 (50% transmission).
The light is backscattered from the retina and traverses the QWP a second time. The circular polarization state (neglecting birefringence of the anterior segment and retina) will be transformed into a linear polarization state that is perpendicular to the incident polarization state. Therefore the light will be reflected at the polarizing beam splitter. The wavefront exiting the pupil plane of the eye will be imaged via all four telescopes onto the DM, the scanners and the SHWS (Haso 32, Imagine Eyes, Orsay, France). Polarization optics is used to minimize influential light on the SHWS caused by backreflections at the lens surfaces. Although all lenses are coated with an antireflection coating specified for the used wavelength region (reflectivity of ~0.4%), residual light intensity arising from backreflections is remaining which would influence the measurement of the SHWS. However, the backreflected light remains in the input polarization state which is not directed to the SHWS (the light is transmitted at the PBS). Additionally a variable aperture stop (500µm to 4mm) is inserted in the focal plane within the first telescope. This plane is conjugated to the retinal plane which allows only light that is backscattered from the retina to pass through. Light backreflected from the lens surfaces and from the cornea is greatly attenuated by the aperture stop. Note that light backreflected from the cornea (after passing twice through the QWP) will be in a polarization state perpendicular to the input polarization state and will therefore be guided to the SHWS. After reflection within the polarizing beam splitter the beam traverses a polarizer (oriented 90° in respect to the first polarizer to eliminate all components with incident polarization state) and is split by a pellicle into two components. A fraction of the light is reflected and directed onto the SHWS while the rest is coupled into a single mode fiber and is detected with an avalanche photo diode (APD module c10508, Hamamatsu, Japan). The signal is recorded with a data acquisition board at 50M samples per second. In this study the system is operated at a frame rate of 10Hz in order to achieve sufficient sampling in y-direction (at least for the small field of views). The images presented were recorded either with 3155 (x) x 1582 (y), 3155 (x) x 791 (y) or 3155 (x) x 395 (y) pixels. The system sensitivity was measured from the signal to noise ratio obtained from a mirror (the beam was attenuated by inserting neutral density filters) with 60dB.
4. In vivo imaging and post processing
In order to evaluate the performance of the system we imaged 5 healthy volunteers. Prior to imaging informed consent from each subject was obtained after the nature and possible risks of the measurement have been explained. The study was performed under a protocol that was approved by the local ethics committee of the Medical University of Vienna and which adhered to the tenets of the Declaration of Helsinki. All subjects underwent a routine clinical eye examination including eye length measurements using a commercial instrument (IOL master, Zeiss Meditec) that is based on partial coherence interferometry . Additionally autorefractometry was performed. Table 1 summarizes the characteristics of the individual volunteers. The mean age of the subjects was 29.6 years. The subject interface of the instrument was a normal head rest (no bite bar). The pupils were artificially dilated; however no drugs were administered to prevent accommodation. The measurements presented from volunteer 5 (contact lens wearer), however, were taken without artificial pupil dilation because after dilation the volunteer suffered from dry eyes which probably caused inferior image quality. Therefore we decided to repeat the measurement at another day without dilation. For this measurement the light in the measurement room was turned off leading to a naturally enlarged pupil diameter during measurement (at 840nm the subjective perception of the light intensity is very low). Individual eye optics corrections (glasses or contact lenses) were not removed for the measurement. Only right eyes were measured in this study although the instrument can be used for both eyes.
A total of 3 seconds were recorded during one measurement. In post-processing the data was first corrected for the sinusoidal motion of the resonant scanner. In a next step, frames that showed severe eye motion (micro-saccades) were removed from the data set. Then all frames were motion corrected using custom software that is based on a cross correlation algorithm. However, due to in-frame distortions a residual motion correction error remains. To eliminate this residual error (similar to a previously published algorithm ) we used custom developed software that splits one frame into several subframes (each containing 100 to 160 horizontal lines) and repeats the correlation for these subframes. Only the central 20 to 40 horizontal lines of the subframe are then used for the corrected image. The next subframe was then shifted 20 to 40 pixels in y-direction and the procedure was repeated until the whole image is retrieved. After this step all frames were averaged to increase the signal to noise ratio. The eye length of each subject was used to convert scanning angles into distances on the retina.
5.1. Small field of view imaging
Figure 3 (Media 1) shows an example of a data set (spanning a field of view of ~1degree) recorded from volunteer 1. Individual cone photoreceptors can be clearly resolved throughout the image. The estimated location of the foveola is marked with an arrow. The averaged frames are displayed in a linear and logarithmic scale (in order to account for the high dynamic range of the image ) in Fig. 4 .Within the central part of the foveola very small cone spacing can be observed in comparison with the surrounding area. However, the regular arrangement is still preserved at this location indicating the ability of the system to resolve foveal cones. In order to quantify the cone spacing we performed FFTs within windows of 50x50µm2 (region of interest ROI) along a diagonal starting from the lower left corner to the upper right corner in Fig. 4 (c.f. white rectangles in Fig. 4(a)) and along a diagonal starting from the upper left corner to the lower right corner. The result is shown in Figs. 5(a) (Media 2) and 5(b) (Media 3), respectively. Within the FFTs a dominating frequency corresponding to the cone row to row spacing can be observed. Close to the fovea instead of Yellot’s rings hexagonal patterns can be observed. At some locations the cone arrangement is extremely regular, therefore in the FFT only the corners of the hexagon can be observed . However, within the central 50µm the corners of the hexagon are visible but less clear. Nevertheless we measured the radius of the hexagon in the central part (foveola) to obtain the row to row spacing of the cones. This distance can be converted into cone to cone spacing and finally into cone density [25,35]. It should be noted that with this measurement we can only provide a rough estimate of the peak cone density, because with this method we can only measure the averaged cone density within the evaluation window. Since the cone spacing exponentially increases with distance from the foveola  we will always slightly underestimate the peak cone density.
With this method we estimated for volunteer 1 a peak cone density of ~210500 cones/mm2. Figure 6 shows images of the fovea recorded from volunteers 2-5. The images recorded from volunteer 2 and 3 are of similar quality than the images recorded from volunteer 1 and foveal cones can be resolved. However, as can be seen in the lower left image foveal cones of volunteer 4 appear less clear. Volunteer 5 was imaged without dilation, therefore the resolution was limited and foveal cones could only be resolved down to an eccentricity of ~0.25°. Nevertheless we estimated the peak cone densities of all volunteers in order to determine the order of magnitude of the cone densities. The estimated peak cone densities of volunteer 2, 3, 4 and 5 were 180500, 124500, 122700 and 154000 cones/mm2, respectively. These values are within the range of data measured with histology  and AO-SLO . For a more accurate determination of the peak cone densities individual cones have to be detected and counted .
5.2. Large field of view imaging
In order to test the performance of the system for larger scanning angles we recorded several data sets of volunteer 1 with increasing scanning angles. Figure 7 shows the foveal cone mosaic recorded with a 2x2 degree field of view. Although the illumination of the patch is not homogeneous the resolution is comparable with the 1x1 degree field of view. Figures 8 , 9 and 10 show the cone mosaic of the same volunteer recorded with a 3x3, 4x4 and 5x5 degree field of view, respectively. Even with the largest field of view (5x5 degrees) individual cone photoreceptors can be resolved. However, at the corners of the image the image quality is degraded and foveal cones cannot be resolved with this large scanning angle. There are basically two reasons for that. First the isoplanatic angle of the eye can be in the order between 3 to 4 degrees  and the SHWS will therefore sense a mean wavefront distortion (averaged over the entire field of view) which degrades AO-correction performance. Second the sampling density in y direction is not sufficient (~1µm pixel to pixel distance) to resolve such small distances (foveal cone spacing is ~2µm). In order to give a better impression of the image quality change associated with the recorded FOV, Fig. 11 shows a comparison of a ROI (50µmx50µm) at an eccentricity of ~0.5° nasal and inferior to the fovea for the different scanning angles. The image quality between 1°x1° and 3°x3° FOV is comparable; however, starting with the 4°x4° the image quality is degraded in comparison to the smaller FOV.
Figure 12 shows a comparison of the image quality in 4 ROI (indicated by rectangles in Fig. 10) in the periphery of the 5°x5° FOV image with the same ROI each recorded with 1°x1° FOV. Clearly, the image contrast of the images recorded with the smaller FOV is higher. Although essentially the same structures can be observed in both images, the separation between individual cones in the images recorded with large FOV is less clear mainly because of residual aberrations (introduced by eye and the instrument) that cannot be corrected for these large scanning angles and because of the lower sampling density.
5.3. Rod imaging
In order to test the ability of the system to image rod photoreceptors we recorded images from volunteer 3 at an eccentricity of ~7° temporal from the fovea. The result is shown on a linear as well as on a logarithmic scale in Fig. 13 . In addition to the cone photoreceptors (c.f. large bright spots in Fig. 13) rod photoreceptors can be resolved (c.f. small regular spots in between the cones).
6. Discussion and conclusion
We presented an alternative approach for AO-SLO using lenses instead of spherical mirrors. Our simulations and the obtained image quality showed, that the performance of an AT based on lenses is comparable to the folded telescope configuration using spherical mirrors. A lens based system could in principle be built very compact. This feature might be of specific interest for commercialization of AO-SLO technology. It should be noted that for our configuration off the shelf achromatic lenses have been used. The performance of a lens based system might be further improved using well designed lenses. On the other hand, one drawback of lens based systems is the sensitivity to chromatic aberrations. However, we think that for the used bandwidth of 50nm the influence of chromatic aberration on the system aberrations can be neglected.
Our system is capable to image foveal cones in 3 out of 5 healthy volunteers. Using FFT’s on the central 50µmx50µm we estimated peak cone densities ranging from 122700 to 210500 cones /mm2 which are in the range of the values from histology and from comparable instruments . As expected from the ZEMAX simulations the image quality of the lens based system is comparable with images retrieved from an instrument using folded reflective telescopes [28,38]. However, a comparison of system performance with other instruments deserves some care. First it should be noted that the DM in our setup consists of only 52 elements which is about half of the elements reported in  and ~1/3 of the elements used in . With our system residual high order aberrations can therefore not be corrected which in the case of volunteers 1-3 did not matter. Second our system utilizes a light source with a central wavelength of 840nm. Therefore the theoretical resolution is slightly degraded in comparison to systems with operating wavelengths around 800nm or 680nm. In this study we did not remove glasses or contact lenses. In fact we found that the AO-performance is slightly better when low order aberrations are corrected using glasses or contact lenses.
Of specific clinical interest are larger scanning angles with AO-SLO (without the need of image stitching). In standard AO-SLO systems (non-folded configuration) the maximum scanning angle is limited by two factors. First the aberrations introduced by the system itself and second, the isoplanatic angle of the individual subject. A lens based system overcomes the first limitation and we could demonstrate that larger scanning angles (up to ~4°) with reasonable image quality can be achieved in healthy volunteers. Similar scanning angles could in principle be achieved using the folded reflective telescope configuration, however comparable investigations have, to the best of our knowledge, not yet been demonstrated. Although at large scanning angles the resolution is slightly degraded it remains sufficient to resolve individual cones down to an eccentricity of ~1° from the fovea (c.f. Figure 9). However, at the corners of the FOV image degradation at larger scanning angles is quite pronounced. We think that the main reason is a focal shift within the images depending on the scanning angle which is introduced by the optics of the eye  and by the imaging system. Although further investigations are necessary, this aberration might be compensated by introducing a scanning angle dependent defocus to the instrument or using dual conjugate adaptive optics .
Figure 13 shows the rod mosaic recorded from volunteer 3 demonstrating the excellent imaging capabilities of the lens based system that are very similar to systems using folded reflective telescopes. Compared to previously reported rod images the field of view of our images is slightly increased which might be beneficial for investigations of the rod mosaic of larger areas.
In conclusion we introduced a new lens based AO-SLO instrument that is capable to resolve foveal cones in 3 out of 5 volunteers and supports larger scanning angles. From the measurements we estimated the peak cone density of 5 healthy volunteers. We further investigated the performance of the system when using larger scanning angles and compared the image quality between data recorded with small and large FOV. Finally we demonstrated the capability of the system to resolve individual rod photoreceptors.
This work was supported by the Austrian Science Fund (FWF project P22329-N20). The authors gratefully acknowledge equipment support from W. Drexler (Medical University of Vienna) and helpful discussions with R. J. Zawadzki (UC-Davis, Sacramento).
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