Kilohertz retinal FF-SS-OCT and flood imaging with hardware-based adaptive optics

A retinal imaging system was designed for full-field (FF) swept-source (SS) optical coherence tomography (OCT) with cellular resolution. The system incorporates a real-time adaptive optics (AO) subsystem and a very high speed CMOS sensor, and is capable of acquiring volumetric images of the retina at rates up to 1 kHz. While digital aberration correction (DAC) is an attractive potential alternative to AO, it has not yet been shown to provide resolution of cones in the fovea, where early detection of functional deficits is most critical. Here we demonstrate that FF-SS-OCT with hardware AO permits resolution of foveal cones, with volume rates adequate to measure light-evoked changes in photoreceptors. With the reference arm blocked, the system can operate as kilohertz AO flood illumination fundus camera with adjustable temporal coherence and is expected to allow measurement of light-evoked changes caused by common path interference in photoreceptor outer segments (OS). In this work, we describe the system’s optical design, characterize its performance, and demonstrate its ability to produce images of the human photoreceptor mosaic.


Introduction
The dashed red line box shows the Zemax™ optical design of the Maxwellian illumination with its beam focused 3 cm away of the cornea. The dashed black line box presents the system operating as a AO-FI camera.
The imaging system operating in AO-FF-SS-OCT configuration consisted of a Mach-Zehnder interferometer with 26 tunable light source (BS-840-2-HP, Superlum, Dublin, Ireland) with 800-875 nm tuning range from 100 nm/s to 27 100,000 nm/s. The light was split into sample and reference arms with a 20/80 beam splitter. In the sample channel, 28 light was focused 3 cm in front of the subjects's cornea, illuminating a 2°field-of-view (FOV) on the retina with a 29 converging (but not focused) beam with a power of 3 mW measured at the cornea, deemed safe for retinal and corneal 30 exposure by the 2014 ANSI standard [3]. The backscattered light by the retina was imaged onto a high-speed 2D 31 CMOS sensor (FASTCAM NOVA S-12, Photron, Tokyo, Japan) allowing frame rates between 1 kHz and 500 kHz 32 over regions of interest between 1024 x 1024 and 128 x 64 pixels. The optical system was designed for imaging 33 through a dilated 6.75 mm pupil, yielding a lateral resolution of 2.6 µm in the eye. To achieve sufficiently high lateral 34 sampling by the camera's 20 µm pixels, a magnification between the retina and sensor of approximately 20x was chosen, 35 accomplished with a series of three 4f telescopes. The sample channel also incorporated a hardware AO subsystem 36 incorporating a superluminescent diode (SLD) beacon (IPSDM0701-C, Inphenix, Livermore CA, USA) with 755 nm 37 center wavelength and full-width-half-maximum (FWHM) bandwidth of 30 nm, a deformable mirror (DM) (DM-97-15, 38 ALPAO, Montbonnot-Saint-Martin, France) and custom-built Shack-Hartmann wavefront sensor (SHWS). The SHWS 39 consisted of a CMOS camera (acA2040-90um, Basler, Ahrensburg, Germany) and microlens array (MLA300-14AR-M, 40 Thorlabs, Newton NJ, USA). Optical power of the beacon was limited to 100 µW to keep the simultaneous illumination 41 from two sources in accordance with the laser safety standards. The AO system was operated in closed-loop at 30 Hz 42 with open source software developed in Python/Cython by our lab [19]. Diffraction-limited imaging (residual wavefront 43 2/16 error RMS ≤ 61 nm -Maréchal criterion) was achieved in all subjects. OCT signal processing was done in Python 44 and MATLAB. 45 By blocking the interferometer reference arm, the setup can also be operated as an AO-FI camera, supporting future 46 studies of light-evoked modulations of common-path interference in the OS. Previous demonstrations have required AO 47 to prevent washout of the responses of neighboring cells due to optical blur, and because in the absense of the reference 48 arm the phase of the backscattered light is not measured, DAC is unlikely to be of use. Previous work also demonstrated 49 that the magnitude of the response was modulated by the coherence length of the illuminating light [23]. Here, the 50 coherence length can be adjusted by altering the sweep range of the source in conjunction with the frame rate of the 51 camera. Further adjustments may be possible in post-processing by incoherently summing frames over known spectral 52 bandwidths. In the present work, in order to ensure all frames had identical spectral content, we required at least one 53 full sweep of 50 nm per frame, which limited us to a maximum frame rate of 2 kHz. Halving the bandwidth would 54 allow us to double the frame rate to 4 kHz, notwithstanding losses in contrast due to shorter exposure times. Still higher 55 imaging rates could be achieved by illuminating with a fixed-spectrum broadband source. (1) Expanding this product reveals the two auto-correlation terms and the two cross-correlation terms: The reference beam was designed to be off-axis relative to the sample arm, by ≈ 1°, in order to create a carrier 62 frequency. By spatially filtering the 2D Fourier transform with respect to the x-and y-coordinates of each camera 63 frame, this approach rejects the DC and auto-correlation components and selects just one of the cross-correlation 64 components [18, 36], resulting in a filtered spectral stack filtered . Doing so accomplishes two goals: first, it removes 65 common path artifacts; second, by selecting only one of the complex conjugate cross-correlation pair, imaging depth 66 is extended [18]. Because the source's sweep speed is constant with respect to , after spatial filtering filtered was 67 resampled to be uniform with respect to wavenumber = 2 / , using linear interpolation, and yielding a -stack of 68 images, filtered ( , , ).

70
In FD-OCT, dispersion mismatch between sample and reference arm is commonly compensated in post processing by 71 adding -dependent phase shifts, defined as a third-order polynomial, to the measured spectral fringe ( , , ) [38]. 72 As with previous reports of slow-sweeping SS-OCT systems [14], the chirp in our measurements was insufficiently 73 corrected by this approach. Higher-order chirp was present, likely due to vibrations causing path length difference 74 variations between reference and sample arms. To correct chirp in our acquired data, a method based on short-time 75 Fourier transformation (STFT) [11,14] was employed. In this approach, the -stack filtered ( , , ) is multiplied by each 76 of a series of boxcar windows of width 8.7 × 10 4 m −1 , centered about the wavenumber corresponding to a frame in the 77 -stack. The windowed -stack was Fourier transformed in the dimension and, in the resulting series of low-resolution 78 volumes, the chirp manifested as apparent axial motion of the sample Δ = − 0 , where 0 is the average sample 79 position in the series. Under these conditions, the phase correction is then given by: which was used to calculate the corrected -stack: Here, the main aspects of the AO-FF-SS-OCT system are presented. We explored the impacts of vibrations on the 83 acquired OCT signal and the importance of STFT dechirp to correct random changes in the optical path length in each 84 acquired volume as well as the advantage of this method with respect to the commonly used polynomial dechirping 85 to improve axial sensitivity. The sensitivity roll-off was also measured. Furthermore, we compared the theoretical 86 and experimentally computed axial displacement sensitivity (or, equivalently, sensitivity to phase shifts) for relative 87 displacement between two scatterers. This parameter will be important to characterize the system's capability to measure 88 ORG responses. The high volume acquisition speed achieved in FF-SS-OCT systems is due to massive parallelization of A-scan 91 acquisition rather than high sweep rates. The relatively slow sweep rate makes the system vulnerable to artifacts 92 caused by system vibrations and motion of the sample. The AO system and associated magnification requirement 93 necessitate two additional telescopes in the sample channel, which may exacerbate these artifacts by increasing the 94 system's footprint.

95
Random changes in the optical path length (OPL) of either arm of the Mach-Zehnder interferometer affect the 96 acquired signal. Their impact depends on their frequency content, and fall into different frequency regimes. Above the 97 sensor frame rate they cause unrecoverable fringe washout. Below the volume acquisition rate (or sweep rate) they 98 result in bulk motion, which can be corrected using registration, histogram-based bulk motion correction [26], or, in 99 phase-based ORG applications, by monitoring relative phase differences within the image [20]. Between these two 100 regimes, however, OPL noise causes time-varying phase shifts within the -stack (chirp) analogous to that caused by 101 dispersion mismatch [38] and, in principle, numerically correctable, as described in §3.1.

102
In point-scanning swept-source systems with sweep rates ≥100 kHz and point-or line-illumination scanning [28] 103 spectrometer-based systems, the high-frequency regime is typically neglected, but the OPL fluctiations between the 104 scan rate and volume rate are encoded as spatially varying phase noise. Parallelization in our system guarantees, in 105 principle, that OPL fluctuations are spatially constant.

106
To characterize system vibrations, we placed a mirror in the retina plane, parked the source at 825 nm, collected 107 images at 1 MHz, and computed the temporal power spectrum of the resulting carrier fringe. System vibrations are 108 illustrated in Fig. 2(a). The power at frequencies above 2 kHz is almost flat, which suggests that little can be gained in 109 fringe contrast in single frames of the -stack by further reducing the single frame integration time. It also suggests 110 that the tradeoff between FOV and volume rate may be unfavorable at volume rates higher than 1 kHz to 2 kHz. The 111 power spectrum shows that most of the power lies at frequencies 1 kHz, which would have negligible effect on the 112 phase stability of typical scanning flying spot SS-OCT systems operating at 100 − 400 kHz sweep rates but would cause 113 significant chirp for swept sources operating at 200Hz -1kHz, and illustrates the importance of a robust method of 114 reducing chirp, either numerically or by increasing the system's speed.

115
To test the effect of increased -stack acquisition speed on the need for numerical dechirping, we acquired series of 116 20 volumes of a mirror placed as a sample, at sweeping rates of 200 Hz, 400 Hz, 1 kHz, and 2 kHz. STFT dechirping 117 was performed to determine the correction phasor ( ) as described in Eq. 3. ( ) was used to visualize the chirp at 118 various volume rates, shown in Fig. 2(b). The blue squares in Fig. 2(a) show the standard deviation of the dechirping 119 phasor ( ) as a function of volume acquisition rate. To facilitate comparison with the power spectrum in Fig. 2(b), 120 we plotted these in log scale. As expected, the standard deviation falls with increasing volume rate, and qualitatively 121 resembles the vibrational power spectrum.  123 To characterize OCT performance, images of a mirror in the sample arm were acquired at 100 kHz (ROI 384 × 240 px), 124 with the source sweeping at 10 000 nm/s, resulting in volume rate of 200 Hz and effective A-scan rate of 18.4 MHz. A 125 neutral density filter (ND=2.0) was placed in front of the mirror, and OCT volumes were acquired. From these, the 126 signal-to-noise ratio (SNR) was estimated by dividing the peak reflectivity of the mirror by the standard deviation of a 127 region in the image that appeared to be free of both signal and coherent artifacts. With a resulting SNR of 14.1, the 128 sensitivity of the system was estimated to be ≈63 dB. Sensitivity roll-off was also measured by axially translating 129 the mirror in the sample arm, using a calibrated translation stage showing an OCT signal drop of 6 dB approximately 130 2.1 mm from the zero path length (Fig. 2(c)).

131
The axial resolution (Δ ) was estimated experimentally by measuring the full width at half-maximum (FWHM) of 132 the point spread function (PSF) with a mirror in the sample arm. Three versions of the corresponding OCT image were 133 computed-without dechirping, with 3rd-order polynomial dechirping, and using the STFT approach. As shown in Fig. 134 3(a), when using the STFT approach, Δ = 10.1 µm in air, close to the theoretically expected by the coherence length of 135 the light source, given by the DC component of the Fourier transform of the spectral interferogram (Eq.2), which was 136 = 9.9 µm. The corresponding resolution in the retina ( = 1.38) was Δ =7.3 µm. As shown in the same figure, the 137 FWHM PSF for the uncorrected and 3rd-order-corrected PSFs were significantly higher. While it is evident that the 138 STFT approach improves the PSF by measuring and correcting chirp, its precision is presumably limited both by finite 139 oversampling of the boxcar-filtered images and shot noise. Thus, residual chirp is likely present, which may have an 140 impact on phase sensitivity, as described in §4.3 below. The intended application of the system is to measure tissue deformations much smaller than the system's axial resolution, 143 which manifest as phase shifts. As reported in early in vivo phase sensitive OCT work [20], when using relative phase 144 differences between two reflective surfaces within the volume, the optical path measurement is immune to axial motion 145 artifacts and other sources of OPL instability.

146
The theoretical lower limit on phase-based displacement sensitivity for a single scattering object is dictated by shot 147 noise, and can be computed from the SNR of the image of that scatterer [8]: For the axial component of the relative displacement between two scatterers, the uncertainty is compounded and 149 given by: where theor, is the displacement sensitivity of the th scatterer.

151
To characterize the displacement sensitivity of our system and the relative contributions of different noise sources, a 152 series of OCT images were collected using a glass cover slip, which contained two bright reflections originating from 153 both surfaces.

154
The SNR of each surface was calculated as described above, and a corresponding value of theor,Δ was calculated 155 using Eq. 6. The reflections originating from the two surfaces of the cover slip had SNRs of 129.0 and 59.6, 156 corresponding to a theoretical displacement sensitivity limit of 7.3 nm.

157
An empirical estimate of displacement sensitivity exp was also made by computing the standard deviation of the 158 phase difference between the two cover glass interfaces, Δ , across 1000 volumes acquired over 5 s: resulting in exp =9.9 nm.

160
This allowed us to estimate the phase noise contributed by the instrument using: The difference between experimentally measured and theoretically predicted displacement sensitivity (Eq. 8) was 162 6.7 nm, which could be due to uncorrected chirp, phase instability in the source, or synchronization between the source 163 and camera, or it may indicate that the OCT system is not shot noise limited.  For retinal imaging with OCT, after -mapping, dechirping, and Fourier transformation, the volumes were segmented 178 axially and the photoreceptor inner segment outer segment junction (IS/OS) and cone outer segment tip (COST) layers 179 were identified, aerially projected, registered, and averaged. Between 10 and 40 projections were averaged to produce 180 each en face OCT image.

181
B-scans from volumetric images of the retina, acquired at 200 Hz, revealed the pair of bright, periodically punctuated 182 and correlated bands in the outer retina, believed to originate from the IS/OS and COST [21]. Unaveraged B-scans are 183 shown in Fig. 4 (c and d). Their appearance was consistent with what is observed in similar scanning AO-OCT images. 184 En face projections of the IS/OS and COST revealed the cone mosaic, as shown in Fig. 4 (e and f), and bear resemblance 185 to AO-SLO images of the cone mosaic, as well as en face projections produced by point-scanning AO-OCT systems. It 186 is notable, however, that due to parallel acquisition of the whole volume, the en face shown here are free from the image 187 warp often visible in images from scanning systems.  189 As described in §2, blocking the reference channel permits the system to be used for high-speed (kHz) AO-FI imaging, 190 with temporal coherence adjustable via the sweep range of the laser and integration time of the camera. This mode of 191 operation permits measurement of stimulus-evoked changes in coherent en face reflectivity of photoreceptors [23], 192 which has been shown to be a quantifiable and reproducible effect [9]. In this mode, the source was set to sweep 193 between 825 nm and 875 nm at sweep rate of 100 000 nm/s.

194
Images of the cone mosaic were successfully collected at 1 kHz (Fig. 5) at 1°and 2°. Cones were laterally resolved 195 in these retinal images. The contrast of the cones is not as high as in comparable AO-SLO images of mosaic. This is 196 presumably due, in parts, to the lack of confocality and the significantly lower integration time. 198 Low order DAC correction has been used to image peripheral photoreceptors using FF-SS-OCT [16]. To estimate the 199 minimal aberration order required to resolve foveal cones, we imaged the the photoreceptor mosaic at 2°T, with the 200 AO correction restricted to certain Zernike coefficient order. This was done by first calculating Zernike coefficients 201 using = + · , where a rectangular matrix containing and partial derivatives of each Zernike polynomial up 202 to the 11th order (75 Zernike polynomials, excluding tip, tilt and piston) at the centers of each subaperture, + is its 203 pseudo-inverse, and is the corresponding set of measured and slopes. Next, coefficients in above the desired 204 order were zeroed (ˆ ) and the filtered slopesˆ were calculated usingˆ = ·ˆ . Mirror commands were generated 205 by multiplyingˆ by the AO system's control matrix.

206
When filtering closed-loop correction by Zernike modes, we observed that residual error increased with a reduction 207 of the number of modes (35-50 nm for = 11 (75 Zernike modes); 70-80 nm for = 7 (33 Zernike modes); 120-140 nm 208 for = 4 (12 Zernike modes); and 950-1020 nm for = 0 (without AO correction). The impact of this residual error on 209 resolution can be seen in the areal images and at the radially averaged power spectra of the images shown in Fig. 6. With 210 7 radial orders or higher a peak can be observed at ≈0.14 µm −1 corresponding to a periodicity of 7.1 µm-within the 211 range of the expected cone spacing at the imaged eccentricity [10,25]. This peak, however, does not appear when ≤ 4. 212 The structures seen on areal images for correction up to n=4 represent random speckle field created by interference 213 between multiple reflectors within the blurred PSF [33]. This result exhibits an advantage of the hardware AO approach 214 with respect to DAC, considering that the latter may be limited by computational tractability [36]. 216 Preliminary evidence suggests that in the ORG, OS elongation velocity increases with increasing stimulus dose [5]. 217 Because OS elongation manifests as phase changes that are wrapped into [0, 2 ], the speed requirements of ORG 218 measurements are dictated, in part, by the desired stimulus levels. We sought to characterize the impact of increasing 219 the OCT volume rate on field of view (FOV), SNR, and displacement sensitivity. To do this, we acquired AO-corrected 220 OCT volumes at rates of 200 Hz, 400 Hz, and 1000 Hz. The IS/OS and COST layers were segmented, and from these 221 SNR IS/OS and SNR COST were calculated as described above. Applying Eq. 6 allowed us to compute theoretical relative 222 displacement sensitivity theor,Δ for each of the volume rates. Representative B-scans acquired at various volume rates 223 are shown in Fig. 7, and resulting values of SNR and theor,Δ are shown in Table 2. The three-dimensional structure of 224 the photoreceptors was visible at all volume acquisition rates, and the displacement sensitivity ranged from 7.3 nm to 225 9.2 nm, sufficient for ORG measurements with bleaching levels as low as 1.8 % [5]. FF-SS-OCT offers an intriguing alternative to traditional raster scanning or flying spot OCT systems, despite some 228 key advantages of these more traditional approaches. In scanning systems, field of view and sampling density can be 229 adjusted by changing scan angles and scanner speed, without any alteration to the optics. Additionally, confocality of 230 fiber-based scanning systems confers an inherently lower noise floor and improved sensitivity and SNR.

231
However, FF-SS-OCT possesses several key advantages as well. Because it requires no pupil conjugate plane(s) 232 for scanners, its optical design is simpler than that of corresponding scanning systems. Moreover, since all parts of 233 the retinal FOV are imaged simultaneously, it is immune to the image warp caused by eye movements in scanning 234 systems. While most high-speed swept-source lasers operate at > 1 µm, slower tunable sources are available at shorter 235 wavelengths, conferring advantages in both axial and lateral resolution. Finally, due to the state of the art in high-speed 236 swept sources and high-speed CMOS sensors, FF-SS-OCT at present confers substantially higher volume rates than 237 scanning systems.

238
In FF-SS-OCT, no light from the sample is rejected by a pinhole or fiber tip when the PSF is degraded by aberrations; 239 it all arrives at the sensor. In flood illumination imaging, this yields little benefit, as the stray light degrades the image 240 SNR. However in the case of OCT, where the phase of the light is measured and can be computationally altered, 241 the possibility exists to reshape the PSF numerically and recover diffraction-limited resolution. Traditional scanning 242 systems may utilize DAC, but in addition to light lost when coupling back also suffer from an ambiguity between 243 aberration and axial eye movement in determining the origins of phase shifts. DAC is an inverse problem requiring 244 substantial computation, and correction sufficient to resolve foveal cones has not yet been demonstrated. However, 245 several investigators have implemented ways to constrain it and improve its tractability [2,12,24]. Nevertheless, the SNR 246 of the OCT image likely imposes a limit on the precision with which we can compute the complex pupil function, and 247 without an analysis of noise propagation in these computations we cannot determine DAC's theoretical resolution limit. 248 The primary application for the system described in this paper is acquisition of optoretinograms (ORGs) of 249 photoreceptors [17] and other retinal neurons [30]. While the ORG is an emerging and rapidly developing tool, at 250 present it requires cellular resolution. In the immediate future we are interested in studying the impacts of retinal disease 251 on photoreceptor function, especially in foveal cones, where such diseases have the greatest impact on quality of life, 252 and rods, which are often impacted first. For ORGs of these, hardware AO confers the required resolution.

253
A related method is FF time-domain (TD) OCT using white light [39], which sidesteps entirely the impact of 254 aberrations on resolution (but not sensitivity) by employing high spatial coherence to filter the photons misplaced by 255 optical aberrations. Phase perturbations manifesting in speckle have been used to visualize subcellular dynamics in 256 tissue explants and the human cornea [35], and may represent a complementary retinal imaging modality, especially 257 when equipped with axial eye tracking [27]. DAC has also been employed in line-scanning TD-OCT, which mitigates 258 the contributions of axial motion to the phase of the interference fringe [13]. In FF-SS-OCT, spatial coherence manifests 259 as cross-talk, and approaches have been demonstrated to mitigate this source of noise [4,7].

260
Another key requirement for current ORG methods is sensitivity to displacements in the retina orders of magnitude 261 smaller than the axial resolution currently offered by OCT. The derivation for displacement sensitivity was described in 262 the context of spectral domain phase microscopy [8], and here we have described how to apply the principle to the ORG, 263 where bulk motion of the retina requires measurement of phase differences instead of absolute phase. We have shown 264 that the system will be able to detect changes of <10 nm in the cone OS, sufficient for detecting responses to stimuli 265 that bleach as small as 1.8 % of cone photopigment. In applications where we would like to detect dysfunction, this 266 sensitivity permits us to detect deviations in the expected amplitude of responses to bright stimuli as small as 5 % cone 267 photopigment bleach level [5], which suggests it could be a useful probe of photoreceptor dysfunction.

268
Displacement sensitivity depends on the SNR of the objects whose displacements need to be measured, and the 269 SNR of these depends on the sensitivity of the system. We found that one of the key limitations to the AO-FF-SS-OCT 270 system's sensitivity was OCT spectral fringe chirp, likely caused by vibrations, in the acquired series of spectral images. 271 We have demonstrated a way to numerically remove this chirp and improve the system's axial PSF and sensitivity using 272 the STFT, although future work includes establishing a theoretical limit for our system's sensitivity and developing 273 additional computational methods to improve it. We have also shown that increasing the system's volume acquisition 274 rate (by reducing its FOV and increasing the camera's frame rate) reduces the amount of chirp present in the spectral 275 fringe. This may be a useful alternative to numerical dechirping when real-time feedback on image quality is required. 276 It may also be a complement to numerical dechirping, but one which sidesteps the bottlenecks imposed by the image 277 SNR and computational tractability. While the fastest retinal volumes we acquired (1 kHz required a very small field of 278 view (∼ 128 µm × 64 µm), this still permits imaging of 300 to 500 cones simultaneously, at eccentricities of 1°to 2°, 279 9/16 enough to study the fundamental properties of the ORG and establish ORG norms. 280 In addition to limiting the FOV, high speed OCT imaging also requires shorter integration times on the sensor and 281 subsequently lower spectral fringe contrast, which reduces the system's sensitivity and its images' SNR. One of the 282 consequences of reduced SNR is poorer displacement sensitivity, and this can be seen in the values computed from 283 retinal images acquired at different volume acquisition rates.

284
The tradeoff between displacement sensitivity and speed may be fortuitous, though. The minimum volume rate 285 required for measuring the ORG is dictated by the initial velocity of the elongation phase. In previous work [5] we 286 observed elongation velocities as high as 3 µm · s −1 , corresponding to 50 rad · s −1 or 8 waves · s −1 , to flashes that 287 bleached 70 % of cone photopigment. These require, minimally, volume rates of 16 Hz to avoid 2 phase wrapping 288 ambiguities, but higher sampling rates offer a better safeguard against them. The initial elongation rates in response to 289 flashes bleaching 1.8 % were approximately 30 times slower, and could be successfully imaged with correspondingly 290 slower volume rates. This means that the responses to dim stimuli, close to the noise floor, can be measured at lower 291 rates with correspondingly higher SNR, while responses to bright stimuli can be measured with higher speeds since the 292 subsequent lower SNR is less of a factor.

293
In this work we have described a retinal imaging system incorporating FF-SS-OCT and hardware AO. The system 294 provides sufficient axial and lateral resolution for resolving foveal cone photoreceptors-which have not yet been 295 visualized using FF-SS-OCT with DAC-and sufficient displacement sensitivity (<10 nm) to measure optoretinographic 296 responses to dim stimuli. The system can double as a flood illumination fundus camera with imaging rates of at least 297 1 kHz. Future plans include incorporation of a DMD-based visible stimulus channel and AO subsystem with higher 298 speed and dynamic range, comparisons of AO and DAC approaches to aberration correction, and investigations of ORG 299 changes in patients with diseases of the retina.  off-axis carrier fringe (thought to be due to system vibrations) measured with a mirror in the sample channel, with the camera fan on and off; blue squares show the standard deviation of chirp ( ( ) ) measured over 20 volumes acquired at selected rates within the frequency range. It is apparent that increasing volume rates reduces chirp, and that this reduction qualitatively follows the power spectrum. (b) (top) Phase correction ( ( )) for 20 volumes acquired with the AO-FF-SS-OCT at different volume rates and calculated numerically using the STFT dechirping approach; the characteristic curve suggests a constant component due to deterministic nonlinearity in the sweep. (bottom) The average ( ) at 2 kHz was used as a proxy for the constant chirp component and subtracted from each phase correction curve to visualize the residual chirp caused by system vibrations or noise in the source. This illustrates the importance of applying high order corrections for -stack dechirping, correcting volumes individually, and the potential advantage of higher volume rates.