Digital micromirror device based ophthalmoscope with concentric circle scanning

: Retinal imaging is demonstrated using a novel scanning light ophthalmoscope based on a digital micromirror device with 810 nm illumination. Concentric circles were used as scan patterns, which facilitated ﬁxation by a human subject for imaging. An annular illumination was implemented in the system to reduce the background caused by corneal reﬂections and thereby to enhance the signal-to-noise ratio. A 1.9-fold increase in the signal-to-noise ratio was found by using an annular illumination aperture compared to a circular illumination aperture, resulting in a 5-fold increase in imaging speed and a better signal-to-noise ratio compared to our previous system. We tested the imaging performance of our system by performing nonmydriatic imaging on two subjects at a speed of 7 Hz with a maximum 20 ◦ (diameter) ﬁeld of view. The images were shot noise limited and clearly show various anatomical features of the retina with high contrast.


Introduction
Retinal imaging is an indispensable tool for diagnosing pathologies associated with the retina. Most retinal pathologies exhibit structural and functional changes in the ocular fundus [1]. Novel methods of observing and monitoring these changes have attracted much attention. The ocular fundus being located in the posterior part of the eye is minimally reflective and moderately scattering [2][3][4] making imaging challenging.
Traditional fundus photography uses so-called flood illumination where a homogeneous broadband light illuminates the whole field-of-view (FOV) [5]. In this type of imaging, both the in-focus and out-of-focus light reflected from the retina is collected by the camera, where the latter reduces the image contrast. To reject the out-of-focus light, pinholes can be implemented to make the system confocal. The confocal arrangement requires scanning using a spot or a line and produces higher contrast images by rejecting out-of-focus light [6][7][8][9][10].
Scanning laser/light ophthalmoscopy (SLO) is used widely in clinical settings for investigation and monitoring of the retina. In a conventional confocal SLO [10][11][12], a single spot is used to scan the retina continuously over two dimensions using scanning mirrors. Significant improvement in speed can be achieved by eliminating one scan dimension and using a line illumination in ophthalmoscopes [13][14][15][16] at the cost of a reduced confocality.
Spatial light modulators such as a digital micromirror device (DMD) have been used for retinal imaging by projecting a series of scan patterns [17][18][19] without using conventional scanning mirrors. The DMD provides temporal and spatial flexibility for illumination which can have several advantages such as optimising the system for speed and confocality. To date, retinal imaging systems based on DMDs either employ a single line scanning method combined with a camera operating in rolling shutter mode [19] or use a large number of parallel lines with a global shutter camera as reported in our prior work [17]. Muller et al. reported acquiring confocal images with a 38 • FOV at 14.3 Hz and the registering and averaging 50 frames into a single high-quality image. Our first generation parallel line scanning ophthalmoscope (PLSO) produced high-quality images with a 10 • × 10 • FOV at an imaging speed of 1.4 Hz although without averaging. At these speeds, image quality can suffer from motion artefacts and a larger FOV is desirable.
In this paper, we present a parallel and circle SLO which uses a DMD to produce concentric circle illumination to fixate the eye and reduce eye motion during imaging. The signal-tonoise ratio (SNR) is improved by eliminating back reflections from the cornea. Corneal back reflections typically have higher intensities compared to the intensity reflected from the retina and reduce the image quality by forming a de-focussed background in the image. An annulus is placed conjugate to the pupil plane to create an annular illumination and the reflected light from the retina is collected through the un-illuminated portion in the centre [20]. A circular aperture in the detection arm blocked most of the corneal reflections but passed the signal from the retina to the camera. The resulting improvement in SNR enabled faster imaging speeds without reducing the image quality.
A model eye was developed to study the SNR improvement obtained with the annulus and a circular aperture with the same outer diameter. The capability of the system for in vivo retinal imaging was demonstrated by imaging two healthy volunteers.

Methods
The study obeyed the principles of the Declaration of Helsinki and was approved by the Institutional Review Board for human use at the VU University medical centre, Amsterdam. A written, informed consent was obtained from each subject. The optical power used for measuring human subjects was in agreement with the maximum permissible exposure limit prescribed by the latest IEC standard 60825-1 [21].

Optical system
The optical layout of the system is shown in Fig. 1A. A light emitting diode (LED) with a bandwidth of 30 nm and a centre wavelength of 810 nm (SFH 4780S, OSRAM GmbH, Munich, Germany) was used for illumination. The half angle divergence of the LED was ±10 • with a maximum radiant intensity of 2900 mW/sr. The use of a LED with low spectral coherence reduced the speckle noise in the final retinal images. The light from the LED was homogenised using a tapered light pipe homogenising rod (Edmund Optics Inc., Barrington, USA) with a square entrance aperture (4 mm × 4 mm). The lenses L1 (f = 40 mm) and L2 (f = 50 mm) relayed the exit aperture (8 mm × 8 mm) of the light pipe onto the DMD.
The SLO used a DMD with a 0.7-inch diagonal micromirror array containing 1024 × 768 micromirror elements with a pitch of 13.6 µm. The Vialux ™ DLP ©Vmodule V4100 board (Vialux GmbH, Chemnitz, Germany) provided control logic and power for the DMD. Each micromirror can be individually programmed to be in an 'ON' or 'OFF' state. A total internal reflection (TIR) prism [22] was used to guide the reflected light from the micromirrors in a perpendicular direction. The 4f system consisting of the lens pairs L3 (f = 37.5 mm), L4 (f = 42.9 mm), the ophthalmic lens L5 (f = 30 mm) and the eye (f = 60 D or 16.7 mm for a typical eye) relayed the light reflected from the DMD 'ON' mirrors onto conjugate positions on the retina. The DMD was magnified by a factor of 0.64 and illuminated the retina with an FOV of 30 • × 21 • . The scattered and reflected light from the retina was imaged on the CMOS camera (ace2040-180kmNIR, Basler, Germany) using lens L6 (f = 40 mm) and the camera objective lens L7 (f = 16 mm). The retinal image was magnified by a factor of 0.72 on the camera forming an image of approximately 1150 × 850 pixels. However, due to the aberrations and field curvature induced by the optics, the maximum area where the retina was in sharp focus corresponded to an FOV of 20 • diameter. Overall, the DMD was magnified by a factor of 0.45 onto the camera plane. Individually, each DMD mirror element was mapped to 1.1 pixels in the camera. The optical paths and beam shapes at different positions in the setup to block the corneal reflections -A point source in the DMD was relayed to a point in the retina using a 4f system consisting of lenses L3-L5 and the eye lens. An annulus was placed at the conjugate pupil plane between lenses L3 and L4 and was imaged onto the cornea using L4 and L5. The reflection from the retina was relayed to the camera using another 4f system consisting of the eye lens and the lenses L5-L7. The reflection from the cornea formed an image again at the conjugate pupil plane in the detection arm and was blocked by the circular aperture, while the retinal signal passed through the circular aperture.
To minimise the stray reflections from the optical elements in the system, we used polarisers P1 and P2 in cross polarised states in combination with the polarising beam splitter (PBS) and the quarter wave plate (QWP). The light from the DMD was linearly polarised to the transmission state of the PBS after passing through P1 and was then circularly polarised after passing through the QWP. The reflection from the retina (and the cornea) again passed through the QWP in the return path, which converted the reflected light to a linear polarisation state aligned with the reflection state of the PBS and passed through P2. In this configuration, the stray reflections by components in the system between the two cross polarisers were rejected.
The camera sensor contained 2048 × 2048 pixels on a 5.5 µm pitch and operated in a global shutter mode. The maximum frame rate at full resolution was 180 frames/second (with 8-bit encoding). The full well charge of each pixel was 11800 electrons (e). A frame grabber (NI PCIe-1473R, National Instruments Inc., USA) was used to acquire and save images. The DMD control board sent the trigger signals to the camera to capture a frame for each pattern. There was a live preview of the image acquired by the camera to make sure that the retina is in focus in the camera.

Annular illumination on the pupil plane and retinal resolution
The diameter of an un-dilated pupil varies according to a lot of factors including age and the luminance of the light used for imaging. The system was designed for a 3 mm pupil diameter as it corresponds to the diameter of an undilated pupil reported in the literature [23]. An annulus with an outer diameter of 4 mm and an inner diameter of 3 mm was made by laser cutting a 1 mm thick sheet of metal. The central obstruction is connected to the outer obstruction by thin metal bridges as shown in Fig. 1B. The annulus was placed conjugate to the pupil plane (A1 in Fig.  1B) between L3 and L4 to create an annular illumination on the cornea with an outer diameter of 2.8 mm and an inner diameter of 2.1 mm. The image of the annular corneal reflections is focussed at the conjugate plane in the detection arm at location A2 in Fig. 1B and has an outer diameter of 3.7 mm and an inner diameter of 2.8 mm. We used a 2.8 mm circular aperture in the detection arm to block the corneal reflections and allow imaging the retina through the central un-illuminated portion of the pupil [20]. Thus, we rejected most of the corneal reflections and reduced the background, thereby improving the SNR. Having a circular aperture of 2.8 mm diameter in the detection arm translates to a 2.1 mm circular aperture on the cornea. According to the Rayleigh criterion, this corresponds to a resolution of 7.9 µm on the retina and 5.6 µm in the camera plane. Since the camera pixel size was 5.5 µm, the resolution was not limited by the camera. In practice the resolution in the retina is worse because of ocular aberrations. For the measurements on the model eye with an aperture, the annulus was replaced by a circular aperture with a diameter of 3 mm.

Parallel scanning method
The DMD can be used to create a large variety of parallel scanning patterns as shown in Fig.  2. The illumination pattern is produced by programming sequential binary patterns which are displaced relative to each other by a constant distance. The parallel scanning strategy has already been showed to provide high-quality confocal images in fluorescence microscopy [24,25]. The parallel line scanning ophthalmoscope (PLSO) [17] has been used to produce in vivo retinal images.
However, there is a natural tendency for the human eye to follow motion, in this case, the moving spot ( Fig. 2A) or line (Fig. 2B) patterns which are projected typically at 100 Hz to 200Hz. This eye movement caused by the subject's inclination to follow the scan patterns leads to motion artefacts in the final images. However, a pattern of concentric circles that shrink provides an illusion of depth where the subject tends to fixate on the centre of the concentric circles where each circle seems to disappear. When the circles are centred in the middle of the DMD (Fig. 2C), the fovea is in the middle of the FOV. When the centre of the circles is shifted to different regions on the DMD, e.g. the edge of the DMD (Fig. 2D), the subject fixates on the centre of these circles bringing different areas around the fovea in the FOV, including the optic nerve head (ONH). Thus, an area around the fovea can be imaged with a scan pattern that at the same time provides a fixation target to the subject. As an example, concentric circle patterns with the centre of the circles at two different locations on the DMD are shown in Visualization 1, Illustrations of the parallel scanning methods used -A: Spots have been used in fluorescence microscopy for imaging stationary biological samples [24]. B: Parallel lines used in our prior work (PLSO) [17]. C: Concentric circles used in our system for scanning with the centres of the circles in the middle of the frame. D: Concentric circle scanning with the centre of the circles at the edge of the frame (see Visualization 1).
which illustrates the tendency to fixate on the centre. Each mirror in the DMD can be programmed individually to be turned 'ON', thereby reflecting the light along the optic axis or turned 'OFF' thus not contributing to the illumination on the retina. Let s be the number of 'ON' mirrors per scanning pattern and p be the total number of DMD mirrors used for illumination. The fill-factor (Γ) for scanning the retina is then defined as (1) N = 1/Γ then gives the total number of patterns required for scanning the FOV. At the end of N measurements, we have an image stack of N frames. By summing all the frames into a single image, we obtain a wide-field image.

Confocal image processing using virtual pinholes
True confocality occurs in an ophthalmoscope which has an illumination pinhole to produce a diffraction limited spot in the retina and a detection pinhole which suppresses the out-of-focus scattered light. Line scanning achieves 'quasi confocality', not all out of focus light is rejected. Some of the scattered light from adjacent illuminated spots along the line dimension can lead to cross talk in the image. This is also true for parallel illumination schemes such as multiple lines or concentric circles. In order to create a confocal image, pinholes should be implemented in the detection arm. These pinholes should be precisely position matched with the illumination pattern to extract just the signal reflected from the sample and reject the out-of-focus light. These detection pinholes are tough to implement as real physical pinholes and hence these pinholes ought to be created in the digital domain using post-processing techniques.
To reduce the scattered light from the retina and suppress the corneal reflections for each pattern in the sequence, we used a post processing method first presented by Heintzmann et al. [26]. The Heintzmann method determines the maximum and minimum grey scale values for every image pixel over all acquired pattern frames (I). The maximum value corresponds to the signal in focus, while the minimum value is a measure of the background signal. The values in between the maximum and minimum relate to different amounts of out-of-focus signal and are therefore excluded from the analysis. The confocal image I con f (x, y) is then calculated as the difference between the maximum and the minimum pixel values as: In this study, the five lowest values per pixel were averaged to obtain a better estimate of the background value for each pixel, and this value was subtracted from the maximum value of that pixel.

Model eye measurements to evaluate SNR improvement
In this section, we describe the SNR model and the experiments done in the model eye to evaluate the SNR improvement with an annulus compared to a circular aperture for different fill-factors. These experiments demonstrate the SNR in the transitional regime between point scanning and full field illumination and form the rationale behind the increase in imaging speed with the annulus without compromising the SNR. A model eye was developed to characterise the imaging performance and to study the SNR improvement when using a circular aperture or an annulus. An uncoated N-BK7 0.5-inch plano-convex lens with a 15 mm focal length (LA1540, Thorlabs GmbH, Germany) and a 7.7 mm radius of curvature was used in the model eye. This radius of curvature matched well to the human cornea anterior radius of curvature of approximately 7.6 mm [27]. The near identical radius of curvature in the model eye ensured that the corneal reflections for the model eye followed the same ray paths as can be expected for the human eye. In addition, the model eye was filled with water to mimic the optical properties of the human eye, which increased the focal length to 20 mm. A uniformly scattering layer of titanium dioxide (TiO 2 ) of 300 µm thickness and scattering coefficient (µ s ) of 10 mm −1 was used as an artificial retina. The amount of power incident on the DMD remained constant as a function of fill-factor, but varied between two values for the annulus and aperture measurement such that the power incident on the cornea was the same for annulus and aperture with the same fill-factor. As a result, the amount of optical power incident on the cornea (180 µW at Γ = 0.05) varied with the fill-factor Γ only for both the annulus and aperture measurements. We determined the readout and dark noise variance for the camera by taking dark frames. The dark plus read-out noise power (σ 2 dark+readout ) was approximately 350 e 2 . We further determined the step size of the A/D conversion by illuminating the sensor with different power levels, establishing the gain factor according to the European Machine Vision Association (EMVA) standard [28] as shown in Fig. 3, which gave a step size of 46 electrons per digital number (DN). The quantization noise then becomes (46) 2 /12 = 176 e 2 . The total measured noise of 350 e 2 minus the quantization noise of 176 e 2 results in a dark noise power of 174 e 2 . Therefore, the shot noise limited regime is reached above 350 detected photoelectrons or a digital number (DN) of 7.5. Our retinal measurements were performed at detection levels of 2500 photoelectrons (DN = 54).
We took the following definition of SNR according to the European Machine Vision Association standard [28]: where η is the wavelength-dependent quantum efficiency of the camera sensor, μ p is the mean photon count, K is the overall system gain and σ d and σ q are the dark noise and quantization noise respectively. In our system, dark noise and quantum noise were negligible for the signal levels we obtain on the camera and hence Eq. (3) becomes, where μ e is the mean number of electrons. Based on the encoding, the number of electrons are converted to a grey value by the camera. These values have contributions both from the light reflected from the retina (retinal signal) and the light reflected from cornea and other surfaces.
If μ e,background is the background and μ e,signal is the total signal on the camera, the light Fig. 3. The measured variance σ 2 sig − σ 2 dark+readout against the mean gray value μ sig − μ dark and the linear regression line used to determine the overall system gain K. σ 2 sig is the variance in signal and μ sig is the mean signal. The graph indicated that the gain of the system is 46 electrons/ DN (1/ slope of fitted line) according to the method described in the EMVA standard [28]. reflected exclusively from the retina is given by μ e,signal − μ e,background . Then the SNR for the retinal signal alone (SN R s ) can be rewritten as [17] SN R s = μ e,signal − μ e,background √ μ e,signal The retinal signal per pixel remains constant for all values of Γ. However, the total background varies linearly with Γ. Eq. (5) can then be written as: with α the retinal signal (μ e,signal − μ e,background ) per pixel and β the background per pixel. Equation (6) was used, as a model to fit the experimental SNR s data. The SNR s for the annulus and the circular aperture as described in section 2.2 was measured for different Γ. When a circular aperture was used the corneal reflections saturated the camera. Reducing the source power below saturation threshold gave a very poor signal with the annulus. Hence a neutral density (ND) filter (OD = 0.6, 35% transmission at 810 nm) was placed in front of the camera when measuring with the circular aperture to prevent saturation. The mean electron count was corrected for this attenuation. The SNR s was estimated for different Γ with the DMD projecting patterns at 140 Hz. For each Γ, one frame (Fig. 4A) was taken, and the signal profile was plotted along different lines as illustrated in Fig. 4B. The mean signal was estimated within a 1/e 2 width of the signal peak and the mean background was measured elsewhere from this profile as shown in the Fig. 4C. The same procedure was repeated several times and averaged.

In vivo retinal imaging
To demonstrate the imaging capabilities of our system, we imaged the right eye of two healthy volunteers. Subject 1 was 32 years old with a refractive error > -1 D and subject 2 was 27 years old with a refractive error > -1.75 D and astigmatism of -2.25 D. Imaging was performed in a dimly lit room with un-dilated pupils (∼ 3 mm in diameter) and with an incident power at the cornea of 180 µW. The maximum permissible radiant power for a collimated static beam at 810 nm wavelength is 639 µW [21]. Laser safety calculations are detailed in the Appendix.
The DMD was used to project concentric circular pattern at 140 Hz with a Γ = 0.05 (N = 20) resulting to an imaging speed of 7 Hz. An annulus with dimensions as described in section 2.2 was used to minimise the corneal reflections. The circular pattern stimulated the subject to fixate on the centre of the circular pattern. Visualization1 provides an impression of the stimulus. The ophthalmic lens could be translated along the optic axis to bring the retina in focus. The measurements were done for 7.2 seconds (1000 frames). Confocal videos were created by updating the image stack with every newly acquired pattern frame and recalculation of the confocal image. As such out of a single dataset, a confocal video with 981 video frames could be obtained. Figure 5A shows the mean signal (retinal signal + background) and mean background plotted for different values of Γ for both the circular aperture and the annulus. When the Γ increases, so does the background and the mean signal. The offset between these two is the retinal signal which remains constant with both illumination methods.  As explained in section 2.2, the corneal reflections which form as a background in the image are substantially reduced with the annulus. The SNR s model (Eq. (6)) was used to fit the SNR s data for different values of Γ. For the circular aperture, the model fit gave α = 1499 e and β = 134645 e. For the annulus, the model fit gave α = 1531 e and β = 20344 e. These values are in agreement with the argument that the retinal signal α remains fairly constant for both circular aperture and annulus for the same power level on the cornea, while there is a 6.6-fold decrease in the background when an annulus is used. This effect translates to increase in SNR s as shown in Fig. 5B.

Model eye measurements to evaluate SNR improvement
When the Γ is as small as possible (for example only one mirror on at a time), the performance of the system falls in line with a conventional single spot scanning SLO. Then the SNR s should be the same with or without annulus, as the corneal reflections would be minimal also in circular illumination case. This effect is already seen in Fig. 5B as the curves start to converge for small values of Γ. But for Γ>0.4, the benefit of using an annulus is pronounced as there is almost a two-fold increase (1.9 at Γ=0.05) in the SNR s . It is also evident from Fig. 5B that a given SNR s can be achieved for a higher fill-factor Γ in the case of an annulus compared to a circular aperture thereby making higher imaging speeds feasible in terms of SNR.
There was still some background when the annulus was used for imaging. This residual background can be attributed to light scattering in the retina which contributes to the background and some small portion of the corneal reflection which reaches the camera. Further, the image formed by a single DMD mirror projected on the retina through an annulus is convolved with the system optics point spread function of the circular aperture in the detection arm to form a slightly blurred spot on the camera. Also, when the corneal reflections are defocused at the circular aperture in the detection arm, they do not get blocked completely. Figure 6 shows the results of in vivo imaging of the foveal region and various peri-and parafoveal regions around the retina of subject 1. Figure 6A is a fundus photograph (Canon CR-2, Canon Europa N.V, The Netherlands) with the coloured circles highlighting the different peri-and parafoveal regions with the fovea in the centre (blue dashed circle). Since we used polarization optics in our system, the macular bow-tie structure is noticeable in Fig. 6D. This structure occurs around the fovea due to a meridional variation in birefringence exhibited by Henle's fibres surrounding the fovea [29]. The different regions (Figs. 6B-6G) were imaged with a 18 • FOV (diameter) by shifting the centre of the concentric circles to different places in the DMD. Figure 7 shows the comparison of an in vivo image with a Spectralis (Heidelberg Engineering GmbH, Heidelberg, Germany) confocal SLO image of the inferior region of the retina of subject 1. 7A is the non-confocal image produced by adding all the individual frames. This image suffers from multiple scattered and out-of-focus light resulting in poor quality with features that are not clearly discernible. The virtual pinholes method was applied to this image in Fig. 7B which shows the resulting high contrast confocal image. Figure 7C shows the average of 5 confocal frames after median filtering with a kernel of 3×3. Figure 7D and Figure   show the single frame and 20 frame average from a Spectralis SLO using the infra red (IR) channel. Figure 7F show the line profiles through the same blood vessels (indicated by numbers 1 through 5 in Fig. 7B) Figure 8A shows the fundus photograph (Topcon TRC 50-DX Type 1A, Topcon Europe Medical B.V., The Netherlands) of subject 2 with a 50 • FOV with the regions imaged with our system denoted by dashed circles. Figure 8B shows the macular region with the fovea (red arrow) in the centre of a 20 • FOV. Figure 8C shows the ONH region with a 20 • FOV (diameter) with blood vessels converging on the ONH. A 'confocal video' with 981 confocal frames is shown in Visualization 2.

Discussion
The ophthalmic imager replaces scanning mirrors and laser sources that are present in the majority of prevailing retinal scanning systems with a DMD and LEDs, potentially reducing cost and increasing flexibility.
The new DMD based ophthalmoscope demonstrates several improvements over our previous system [17]. The background from corneal reflections was significantly reduced with the use of annulus, which increased the SNR by almost a factor of two. Moreover, the reduced corneal background made it possible to use higher fill-factors in the DMD, thus improving the imaging speed of the system from 1.4 Hz to 7 Hz without any loss in SNR. Although the concept of annular illumination has been well known for many years, this paper offers a quantitative analysis of the benefit of such illumination when going from the low fill-factor regime (conventional confocal SLO) towards flood illumination. The images obtained with our system are comparable to images obtained with a commercial SLO. Shifting the centre of the concentric circles to image different regions is a novel way to provide fixation while imaging different periand parafoveal regions.
The system is modular with respect to the light source and enables changing the wavelength for imaging. Adding new wavelengths will make the system achieve multi-spectral imaging. Multi-spectral retinal imaging has gained interest in recent years [30,31] as it helps gathering quantitative information on retinal health.
The in vivo images sometimes show faint circular artefacts after post processing. This is attributed to motion artifacts creating a discontinuity in scanning. The 'confocal video' (see Visualization 2) shows the eye motion occurring while imaging over a longer period. It could be possible to extract motion information and use this to track the eye's movement. When done fast enough, it will also enable doing eye motion correction in real-time [32,33] to improve the quality of the images by averaging multiple images at the same location and correct for the minor motion artefacts present in these images. Image registration could potentially be better with the parallel illumination as all the illumination points of a single frame are undistorted.
The primary shortcoming of the system is that the illumination light efficiency is poor. The DMD is operated at low fill-factor resulting in the majority of power being unused. Hence a high power and low divergence LED is needed to provide sufficient power for imaging the retina. Although a low divergence LED was used in our system, this might not be available for all wavelengths, particularly for the wavelengths of interest for multispectral imaging. This is expected to improve in the future with the progress in LED source development. As we currently are a factor of 4 below the maximum permissible exposure according to the IEC standard 60825-1 [21], there is still considerable room for further improvement. The polarisation optics to reduce the stray reflections in the system lead to polarisation artefacts in the images due to birefringence of the cornea [34] and retina [29], influencing the images being recorded The position of the cornea in the imaging system is critical for the annulus to provide complete rejection of the corneal reflections. If the cornea is not positioned in the optimal imaging plane, the corneal reflections are not totally blocked, and hence they reach the camera. Although the subject's head is placed on a chin rest, small movement can occur. A pupil camera which tracks the location of the beam on the cornea can be implemented, allowing the repositioning of the illuminating beam to the centre of the cornea.

Conclusion
We have developed an SLO using a DMD displaying concentric circle pattern to scan the retina and provide a fixation target to the subject to fixate the eye. By changing the centre of the concentric circle pattern, different regions could be imaged. We used an annulus in the illumination arm to create an annular illumination on the cornea and block most of the corneal reflections using a circular aperture in the detection arm. Using the annulus there was a 6.6-fold decrease in the background translating to a 1.9-fold increase in the SNR. This translates to a nearly 7-fold increase in speed for Γ= 0.05 at the same SNR. We demonstrated the experimental implementation of the imaging system by imaging two subjects in vivo with a maximum FOV of 20 • diameter at a 140 Hz sub-frame rate and 7 Hz full frame rate.

Appendix: Laser safety calculations
Calculations for the maximum permissible exposure (MPE) on the retina were done using the latest laser safety standard IEC 60825-1 published in 2014 [21]. Although we used concentric circle illumination on the retina which extends to the whole field of view, the most conservative estimate of the MPE can be obtained by assuming a collimated static beam entering the eye. For the limiting aperture, a pupil size of 7 mm was used. From the IEC standard, the MPE expressed as irradiance or radiant exposure for wavelength range 700 nm to 1050 nm for a collimated beam (C 6 = 1) with exposure time of 10 seconds or longer is given by where C 4 and C 7 are correction factors and A Aperture is the pupil area. The equations for the correction factors can be found the Table 9 of the standard. The C 4 and C 7 are both dependent on the wavelength and with 810 nm illumination the C 7 equals 1 and C 4 can be calculated as: C 4 = 10 0.002(λ −700) = 10 0.002(810−700) ≈ 1.66 (8) where λ is the central wavelength of the illumination used. After adding the correction factors and the area of the limiting aperture, the maximum permissible exposure limit for 810 nm illumination can be expressed as However, the concentric circles are projected within the whole field of view and the power is distributed amongst all of the circles, which makes this MPE estimation very conservative. If the illumination of the retina is assumed to be what the standard calls extended view, the MPE limit will be more than an order of magnitude higher.