For more than a decade, two types of raster-scanned retinal imaging systems equipped with adaptive optics (AO) have been used to study the structure and function of the living human retina. The first of these is AO scanning light ophthalmoscopy (AO-SLO) [1], in which light scattered by the retina is measured using a detector, while using a pinhole conjugated to the retina in order to reject out-of-focus and multiply scattered light. AO-SLO produces a two-dimensional, areal image of the retina with cellular resolution. It has been used to image the cone mosaic [1] and rod mosaic [2] and, with modifications to the pinhole size and position and/or numbers of detectors, it has been used to image retinal vasculature in detail [3], cone inner segments [4], and inner retinal neurons [5]. The second raster-scanned AO imaging modality is optical coherence tomography (OCT), which permits volumetric imaging of the retina [6], including 3D images of single cells [7] when combined with AO [8,9]. Each of these modalities offers unique advantages, with SLO allowing detection of multiply scattered photons, as well as fluorescently emitted ones and OCT allowing extraction of axial scattering profiles and sensitivity to sub-wavelength changes in the axial position of scattering structures. Because of its relatively small data size, real-time visualization of the AO-SLO image stream is computationally practical, whereas the large data size, dimensionality, and post-processing requirements hamper real-time visualization of the AO-OCT stream. Incorporating an SLO channel in the system permits visual validation of AO performance and rapid tuning of optical, electronic, and software parameters. The tractability of AO-SLO data also permits imaging and montaging over large fields of view (FOVs). On the other hand, the AO-OCT provides a three-dimensional volume containing information about the amplitude and phase of the reflected light, the latter of which has been used effectively to measure blood flow [10], improve vascular contrast [11], and measure physiological changes in photoreceptors [12].

In the years since the emergence of AO-OCT [8,9], a number of investigators have integrated SLO detection into AO-OCT systems [13–20]. However, due to the comparatively slow acquisition speeds of volumetric OCT images, parallel acquisition of coextensive OCT volumes and SLO frames has not yet been successfully achieved. From the outset, AO-SLO frames have been acquired at rates near 30 Hz [1] with sampling rates of tens of megahertz. Acquisition speeds of OCT systems depend on the speed and sensitivity of the linescan camera (in spectrometer-based systems) or the sweep rate of the tunable laser (in swept-source systems). These characteristics are continually improving, and the recent advent of Fourier-domain mode-locked (FDML) swept sources with sweep rates greater than 1 MHz [21,22] bring the sampling frequency of OCT significantly closer to that of SLO. In this Letter, we introduce a combined AO-SLO-OCT system that leverages a 1.6 MHz FDML laser in order to acquire synchronized, coextensive OCT volumes and SLO frames. This mode of acquisition is significant because it broadens the scientific capabilities of the system by enabling, for example, simultaneous measurement of light-evoked OCT phase changes and SLO scattering changes. Moreover, it simplifies the system’s optical design and data processing by obviating separate scanning channels for OCT and SLO and producing corresponding images, respectively. The scientific benefits and practicality of our combined approach will grow with the speed of future swept sources.

A schematic of the sample arm in the combined AO-SLO-OCT system is shown in Fig. 1A. The swept-source OCT system employed an FDML laser (FDM-1060-750-4B-APC, OptoRes GmbH, Munich, Germany) operating at an A-scan rate of 1.64 MHz [21,22]. The details of the OCT system can be found in Ref. [23]. In summary, a Michelson interferometer with three 50:50 fiber couplers was used to split the light between the sample and reference arm, and the interference pattern was detected using a balanced detector, with a measured sensitivity of $-85\mathrm{dB}$. A dichroic mirror (zt1064rdc-sp, Chroma, U.S.) was employed to reflect the OCT beam while transmitting the SLO light. The scanning system contained a resonant scanner (SC-30, Electro-Optical Products Corp., Ridgewood, NY, U.S.) oscillating at 2 kHz in the horizontal direction and a galvanometer scanner in the slower vertical direction. The scanner configuration, in conjunction with the A-scan rate, permitted the acquisition of six volumes per second over a FOV of $1°\times 1°$ (400 A-scans per B-scan and 340 B-scans per each volume). Because the laser has a fixed A-scan rate, a trade-off exists between spatial sampling density, FOV, and volume rate. An earlier version of the AO-OCT [23] employed a 5 kHz resonant scanner to achieve a 30 Hz volume rate, but with a smaller FOV and lower density; future work with the present system requires this larger FOV and density, and will not be limited by the current 6 Hz volume rate. SLO images were acquired by focusing light from a separate 840 nm SLD source with a 50 nm full width at half-maximum bandwidth (FWHM) (S840, Superlum Ltd., Moscow, Russia) on the retina. A wedged beam splitter was employed in front of the SLO fiber collimator to reflect 10 percent of the light toward the sample arm (BSN11, Thorlabs, U.S.). In the detection channel, the back-scattered light from the retina was spatially filtered using a confocal pinhole with a diameter of 0.5 Airy disk and was detected by a photomultiplier tube (PMT) (H7422P-50, Hamamatsu, Japan). The output of the PMT module was amplified using a high-speed current amplifier (HCA-10M-100K, Femto, Germany) and digitized by a frame grabber (ATS9440, AlazarTech, Canada). Table 1 summarizes the characteristics of both SLO and OCT systems and their data acquisition settings during imaging. The powers of the OCT and SLO beams were 1.8 mW and 150 μW at the cornea, respectively. Together, they are less than 50% of the maximum permissible exposure (MPE) for eight hours of continuous viewing, as specified by ANSI, with the SLO power being 8% of MPE and the OCT power being 36% of MPE [24].

 

Fig. 1. (A) Sample arm of the combined AO scanning system: DM, deformable mirror; SHWS, Shack–Hartmann wavefront sensor; AL, achromatic lens; S, spherical mirror; FM, flat mirror; BS, beam splitter; DBS, dichroic beam splitter; HS, horizontal scanner; VS, vertical scanner; BD, beam dump. Sub-pixel strip-based registration of the cone mosaics at 1.75°TR in a normal subject: (B) and (C) show an average of 50 motion-corrected cone mosaics acquired 1.75° temporal to the foveal center simultaneously by SLO and OCT systems, respectively. (The scale bar is 50 μm across.)

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Table 1. Specifications of the Combined AO SLO/OCT System and Scanning Parameters During Imaging

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The data acquisition platforms for SLO and OCT were on two different computers and, thus, required synchronization of the channels during data recording. For this purpose, the OCT B-scan trigger signal was shared between the two digitizers and the OCT/SLO systems operated in master/slave configuration. During a one-way trip of the resonant scanner, the OCT digitizer captured a B-scan of 400 samples, while the 20 MHz SLO digitizer acquired a line scan of 4096 samples. This design permitted simultaneous acquisition for SLO and OCT channels (see Visualization 1).

The AO subsystem incorporated a high-speed deformable mirror (DM-97-15, ALPAO SAS; Montbonnot-Saint-Martin, France) and a Shack–Hartmann wavefront sensor (SHWS) consisting of a $20\times 20$ lenslet array (diameter 10 mm, pitch 500 μm, $f=30\mathrm{mm}$; Northrop-Grumman Corp, Arlington, VA, U.S.A.) in front of a scientific camera (Ace acA2040-180 km; Basler AG). In this design, the SLO light also served as a SHWS beacon, where 8% of the light coming from the eye was reflected toward the SHWS using a beam splitter. The system measured and corrected aberrations over a 6.75 mm pupil in a closed loop at a rate of 10 Hz, yielding a theoretical diffraction-limited lateral resolution of 2.75 μm and 3.2 μm for the SLO and OCT channels, respectively. Custom software developed in Python/Cython controlled the AO loop [25]. After obtaining informed consent, two subjects free of known retinal disease were imaged. Each subject’s eye was dilated and cyclopleged by instilling topical drops of 2.5% phenylephrine and 1% tropicamide. All procedures were in accordance with the tenets of the Declaration of Helsinki and were approved by the University of California, Davis Institutional Review Board. A bite-bar and forehead rest were used with a motorized $X-Y-Z$ translation stage to position and stabilize the subject’s pupil during imaging. To position the eye at specified retinal locations and also to reduce eye movements, a calibrated fixation target was employed during imaging. A dichroic mirror was placed in front of the pupil to combine the imaging with the fixation target beams. The average time for acquiring one data set, including the initial alignment, was less than a minute with more than 2 min between the subsequent acquisitions. The post-processing of the acquired OCT data was done using custom software [26] developed in Matlab (MathWorks, Inc., Natick, MA, U.S.), as was image segmentation and visualization. First, the volumetric images were segmented axially, and the inner-outer segment and cone outer segment tips layers were automatically identified and projected by summing over their axial extents. A calibration grid target (R1L3S3P, Thorlabs, NJ, U.S.) was imaged prior to the experiments to compute and compensate for the sinusoidal distortion of the images due to the resonant scanner. A single projection was selected as a reference image, and the remaining target images were divided into strips of between three and 15 pixels of height, and registered to the reference using a strip-based registration approach [12,27,28].

In Fig. 1, panels (B) and (C) show the motion-corrected average of 50 cone mosaic images from the foveal center and 1.75° temporal to the foveal center (TR) acquired by SLO and OCT channels, respectively. Despite being able to visualize cones at 1.75° with both AO-SLO and AO-OCT, the cones in the OCT en face image, while having the same locations and density as those in the SLO image, appear larger in diameter. Their apparent size may also be affected by their waveguiding properties. In single mode fibers mode field diameter (MFD) increases with increasing wavelength. Foveal and parafoveal cones are thought to behave like single-mode fibers, so the difference in wavelength between the two channels may cause a difference in their apparent size [29,30]. As shown in Fig. 2, panels (A) and (B), the foveal cones are clearly visible in the AO-SLO image, but not in the AO-OCT image. Part of the apparent difference in foveal cone visibility is due to a wavelength-dependent difference in lateral resolution. It is also possible that coherent interference between cones, analogous to speckle noise, contributes to their reduced visibility in the OCT image [31].

 

Fig. 2. (A) and (B) show an average of 50 motion-corrected cone mosaics at Fovea from the SLO and OCT channels, respectively. (The scale bar is 50 μm across), (C) and (D) show the average of 50 motion-corrected cone mosaics at 3.5° TR from the SLO and OCT channels, respectively. (E) and (F) a small FOV of the SLO frame and OCT en face image at 3.5° TR, color-coded (SLO in green and OCT in magenta) and superimposed on top of each other before and after rigid-body registration, respectively. In (F), the overlapped area after rigid-body registration appears more in white. (F) also shows the 2D cross-correlation map of the SLO and OCT images. The observed shift could be the the result of beam misalignment in the image plane or transverse chromatic aberration, possibly exacerbated by beam misalignment in the pupil plane.

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Multimodal systems with multiple imaging wavelengths suffer from longitudinal and transverse chromatic aberration (LCA and TCA, respectively). LCA is due to wavelength-dependence of focal length and can be compensated for by adding defocus to one of the beams—here the OCT beam. In TCA, beams with different wavelengths focus in the same plane but with a lateral offset. While LCA is relatively constant for different imaging eccentricities, it has been shown that TCA varies for different eccentricities as the angle between the imaging beams and optical axis changes [32–34]. As shown in Ref. [34], beam position in the pupil can also impact TCA. To show the offset in our system, a small FOV of SLO and OCT images in Figs. 2C and 2D were color-coded and superimposed on top of each other before and after rigid-body registration and are shown in panels (E) and (F), respectively. The 2D cross-correlation between the SLO and OCT channels indicates an offset of 19 μm between the channels. It has been shown that TCA can result in beam displacements of several arcmin in the parafovea [33] over 300 nm in the visible and NIR, which is roughly consistent with our finding. Additional potential sources of image misalignment are electronic latency between frame grabbers and misalignment of the beams in the pupil plane, the latter of which can exacerbate TCA. The displacement of spots on the retina is tantamount to a temporal delay between the images. In the worst case, when the displacement is in the slow (vertical) direction, 19 μm corresponds to approximately 10 ms of delay. In the images shown in Fig. 2, the delay was approximately 6 ms. This temporal delay or spatial offset may limit the precision of measurements requiring absolute synchronization; therefore, real-time TCA measurement and compensation may be required [34].

The SLO-OCT images acquired at 6° TR for the same subject are shown in Fig. 3. Single and motion-corrected average of 50 OCT en face projections are shown in panels (A) and (B), respectively. The corresponding single and averaged SLO frames are shown in panels (C) and (D), respectively. A small region in both averaged images was magnified for better visualization of rod and cone photoreceptors. For a detailed comparison between the spatial morphology of the photoreceptors in the AO-SLO and AO-OCT images, panels (B) and (D) were color-coded (SLO in magenta and OCT in green) and shown in a pseudocolor composite in panel (E). As demonstrated by panel (E), there is correspondence between most of the cones, but not for rods. We believe that the difference in lateral resolution, as well as coherent crosstalk (described above) may impede the visualization of complete rod moasic with the OCT.

 

Fig. 3. (A) Single and (B) motion-corrected average of 50 OCT en face projections (logarithmic scale) acquired at 6° TR. The lower panel shows the (C) single and (D) motion-corrected average of 50 SLO frames (logarithmic scale). (E) Superimposed color-coded OCT (green) and SLO (magenta) cone mosaics from (B) and (D). (F) Magnified image of the marked area on (E) which shows the spatial correlation between rods and cones from the two imaging systems; where the magenta and green channels overlap, the image appears white. The scale bar is 50 μm.

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In this Letter, we have described the results of a combined AO-SLO-OCT with coextensive, synchronized imaging. Synchronized imaging provides an avenue for a number of interesting experiments, but the precision of those measurements will be limited by TCA and the alignment of the two beams.