Integrated fixation and stimulus channel for adaptive optics ophthalmoscopy

Marcel T Bernucci; Yan Liu; Qiuzhi Ji; James A Crowell; Kristen Bowles-Johnson; Matthew J Keller; Donald T Miller

doi:10.1364/BOE.574539

. 2025 Oct 31;16(11):4796–4820. doi: 10.1364/BOE.574539

Integrated fixation and stimulus channel for adaptive optics ophthalmoscopy

Marcel T Bernucci ^1,^†,^*, Yan Liu ^1,^†, Qiuzhi Ji ¹, James A Crowell ¹, Kristen Bowles-Johnson ¹, Matthew J Keller ¹, Donald T Miller ¹

PMCID: PMC12643021 PMID: 41293711

Abstract

Adaptive optics (AO) ophthalmoscopes allow high-resolution imaging of retinal structure and function at the cellular level. Due to their high magnification and small field-of-view (FOV), these systems require precise fixation and light delivery to control the retinal region being imaged and stimulated. We present a high-efficiency fixation and stimulus channel for AO ophthalmoscopy, offering an extended working distance, wide steering range, and broad dioptric correction. For stimulation, the channel delivers intense, near-monochromatic light flashes across much of the visible spectrum. Our design uses all stock components, except for a 3D-printed conic mount and a few machined parts. We balance key system trade-offs and demonstrate design performance through several AO optical coherence tomography (AO-OCT) structural and functional imaging examples. Although originally developed for the Indiana AO-OCT system, these design principles can be readily applied to other AO ophthalmoscopic platforms.

1. Introduction

Adaptive optics (AO) ophthalmoscopy continues to undergo significant advances, improving both the measurement and correction of wavefront aberrations in the living human eye [1–9] and enabling better imaging of retinal structure and function at the cellular level [8,10–22]. Like conventional microscopes, AO ophthalmoscopes provide high magnification over a limited field of view (FOV). However, unlike microscopes – where specimen positioning is controlled via a precise XYZ translation stage – AO ophthalmoscopes depend on the subject’s ability to direct their gaze. Therefore, to control the retinal region being imaged, AO ophthalmoscopes include an auxiliary optical channel that projects a visible fixation target, typically computer controlled. The accuracy and stability of the fixation control have become increasingly important as AO ophthalmoscopes are more frequently used in longitudinal studies requiring repeated imaging of the same retinal areas over time.

Despite the technical challenge of conforming the fixation channel design to the stringent requirements of AO retinal imaging and the central role fixation control plays in enabling reliable and repeatable AO retinal imaging, the fixation channel remains one of the least described components of AO ophthalmoscopic systems. Few details are provided in the literature, and to the best of our knowledge, only one publication by Steven et al. [23] has been dedicated to this topic. In Steven et al.’s report, they describe a fixation channel design utilizing a combination of custom and off-the-shelf optics for AO scanning laser ophthalmoscopy (AO-SLO) imaging to achieve wide field viewing of 32.2° x 32.2°, a working distance of ∼45 mm (see Discussion for details), and a large dioptric correction range of -25 to +20 diopters by mechanically translating an integrated Badal optometer. While highly precise and effective, this configuration requires translation over nearly 2 feet for refractive error correction, resulting in a large system footprint and total unfolded optical path length of ∼ 1 m. Despite custom optics and the system’s size, at least one other group replicated Steven et al.’s design that we know of [24], underscoring its appeal.

Ferguson et al. [25] describe an alternative approach for eccentric AO-SLO retinal imaging that uses two large spherical mirrors to dynamically steer the AO ophthalmoscope imaging beam around the subject’s fixation point, enabling imaging across ±15° (30° total). By offsetting the fixation target up to ±15° relative to the foveal center, imaging at even more peripheral retinal locations is possible ([26], personal correspondence). However, the use of large steering mirrors reduces the design’s compactness and complicates integration with existing AO ophthalmoscopes. Additionally, the large mirrors introduce systematic astigmatism and other field-dependent aberrations that vary with retinal eccentricity, necessitating additional AO correction and reducing the available deformable mirror stroke to correct ocular aberrations for diffraction-limited imaging. While many other groups have demonstrated AO imaging in the peripheral retina utilizing various fixation channel configurations, the implementation of these systems is widely underreported and their design principles remain unclear. Given this lack of detailed reporting and the limitation of existing fixation channel designs, there remains a strong need for a compact, low-cost, and easily integrable fixation channel design that provides an extended working distance, wide steering range, and broad dioptric correction for AO retinal imaging.

In addition, many modern AO ophthalmoscopes incorporate an auxiliary visible stimulus channel to perform optoretinography (ORG)—the measurement of optical changes in the retina in response to light [14,17,19,20,27–35]. Because both fixation and stimulus channels operate in the visible spectrum, the design requirements of the stimulus channel impose further limitations on the fixation channel. For the stimulus channel, it is desirable to deliver retinal stimuli across the visible spectrum, with a wide power range, and short (millisecond-scale) flash durations. Additionally, achieving near-monochromatic stimulation (narrow bandwidth) requires suppression of other wavelengths. Broadband light sources such as LEDs and supercontinuum lasers inherently emit over a wide spectral range and exhibit sidelobes (unwanted energy at wavelength adjacent to the stimulus wavelength of interest). Without sufficient spectral filtering, these stray wavelengths can contaminate the functional responses of retinal cells, confounding the association between stimulus wavelength and the excitation. This is particularly true for S, M, and L cone photoreceptors that respond across more than six orders of magnitude of light intensity [36]. Although slightly broader spectra and significant spectral tails can suffice for applications such as cone classification [17], narrowband stimuli provide greater spectral precision and robustness in characterizing functional retinal biomarkers such as cone spectral sensitivity. While various stimulus channels have been implemented successfully, none to our knowledge meet all these performance requirements and most systems are restricted to one or two excitation wavelengths [14,16,19,27,29–31,37,38]. Stimulus systems supporting multiple wavelengths enable more robust and accurate classification of S, M, and L cone spectral types [12,17,20,24,32,35,39,40], and allow for additional measurements, such as cone spectral sensitivity [35], photopigment density variation [35], and functional changes in cone photoreceptors in diseased eyes [20].

It is in this context that we developed an integrated fixation-stimulus channel designed to address many of these challenges and evaluated its performance within the Indiana AO-OCT system. For the fixation subsystem, we balanced the trade-offs between subject working distance, steering field of view, and dioptric correction range. For the stimulus subsystem, we prioritized performance in terms of wavelength control, exposure duration, power level, and beam uniformity at the retina. With the exception of a 3D printed conic optical mount and a few generic machined aluminum parts, our system was built from stock components to reduce cost and lead time. This resulted in our use of two inexpensive high-end commercial lenses (an aspheric Volk Optical lens and a 7-element Canon camera lens). While these high-end lenses enhanced system performance, their proprietary designs limited our ability to perform theoretical system evaluation using commercial optical ray tracing software. Consequently, we relied on laboratory measurements for performance validation.

Here, we present the objectives, design considerations, and performance of our integrated fixation-stimulus channel, supported by AO-OCT imaging examples that demonstrate its utility for investigating retinal structure and function at the cellular level. Although our fixation-stimulus channel was developed specifically for the Indiana AO-OCT system [41], the underlying design principles, component choices, and performance trade-offs are broadly applicable to other AO ophthalmoscopic platforms.

2. Methods

The design of the fixation-stimulus channel (Fig. 1) was guided by several overarching objectives, which can be organized by subsystem. We begin with the objectives and design of the fixation subsystem, followed by those of the stimulus subsystem and its integration with the fixation subsystem and the AO-OCT retinal imaging system. We then describe several experiments conducted to validate the performance of our fixation-stimulus channel, along with the AO-OCT imaging and optoretinography stimulation protocols employed. Detailed 3D renderings, parts drawings of the integrated channel, and a comprehensive system parts list are provided in the Supplement 1^{(3.3MB, pdf)}.

2.1. Fixation subsystem design

The fixation subsystem design was driven by four main performance objectives: 1) to direct the subject’s gaze >18° retinal eccentricity (in any direction and beyond the nasal rim of the optic disc) in order to cover the fundus area most critical for monitoring disease, 2) to correct a 25 D range of spherical refractive errors covering those most commonly found in the population, 3) to provide a large working distance between the fixation subsystem and the subject’s eye to accommodate individuals with varying head anatomies (e.g., deep-set eyes) and facial hair, and 4) to integrate the fixation subsystem with the stimulus subsystem.

We found that a cascade of two lens-based relay telescopes, with the final one in a Badal configuration using the eye’s optics as the last element, provided a simple layout with sufficient degrees of freedom to meet these objectives. The first telescope was designed to integrate the stimulus beam (discussed later) and axially translate to correct the subject’s refractive error. Two high NA (numerical aperture) lenses (L1 and L3) provided a large FOV fixation steering range, while providing sufficient working distance to accommodate subjects with deep-set eyes and facial hair. We limited the diameter of all optics to ∼2” to reduce the bulkiness of the subsystem and expand the selection of commercially available lenses. These aspects of the subsystem design and the various components used to realize them are detailed below.

The fixation subsystem starts with an active-matrix organic light-emitting diode (OLED) microdisplay (DSVGA150 White OLED-XL, eMagin) following Liu et al. [24]. The OLED presents a white, computer-controlled, fixation target (a cross symbol in our case) on a black background, which is viewable by the subject through the two relay telescopes. We chose the compact, high performance Bild BIT 1011A1 driver board because the OLED is mounted directly at the board’s geometric center. This facilitated attachment to a cage system and allowed XY and rotational alignment of the display along the optical axis of the fixation subsystem. The OLED’s “Deep Black” architecture rendered off-pixels as true black, achieving a contrast ratio of 10,000:1. This minimized unintended excitation of cone photoreceptors – an important consideration for cone ORG measurements during imaging. The microdisplay has 800 × 600 pixels with a pixel pitch of 15 µm (nominally 17.5 µm at the retina after imaging through the relay telescopes). The display brightness is specified at 900 cd/m² (typical) and up to 2000 cd/m² (maximum), more than an order of magnitude brighter than what we empirically determined was needed for reliable viewing of the fixation target. We took advantage of this high brightness by using a 10:90 (Transmit (T): Reflect (R)) beamsplitter (BSX16, Thorlabs) to inject the stimulus beam (described later) into the fixation subsystem. This increased the light throughput efficiency of our stimulus beam by 1.8× compared to using a 50:50 beamsplitter, while reducing the microdisplay brightness by 90%.

The microdisplay was placed at the front focal plane of L1 (Canon, f_eff = 50 mm, f/1.4, 7-element USM lens; entrance pupil diameter = 35.7 mm) and its image was relayed to the eye using L2 and L3. The Canon lens features a modified double-Gauss optics design to control distortion and spherical aberrations. L2 consists of a pair of inverted achromatic doublet lenses (2 × ACT 508-500-A, Thorlabs, f_eff = 250 mm, 2” mm diameter), a common lens design strategy to reduce chromatic aberrations and field curvature in infinite conjugate systems [45], and in our case across the projected field of the microdisplay. This design principle has also been implemented for scanning systems using visible light (e.g., visible light OCT) [46].

Lens L3 is an aspheric singlet (Volk, f_eff = 71.4 mm (14 D), entrance pupil diameter = 52 mm), selected for its thin profile to fit into our custom aluminum conic mount. It is designed for high image quality (particularly low distortion) for retinal viewing in binocular indirect ophthalmoscopy. Importantly for this application, its relatively long focal length provides a large working distance (WD) of 27 mm and 42 mm, measured from the top edge and center of the dichroic mirror (DM), respectively, to the corneal plane (Fig. 1). The conic mount was essential not only for housing L3 but for simultaneously coupling the fixation-stimulus channel to our AO-OCT imaging system via a low loss, low wavefront error, and low group delay dispersion dichroic mirror (Thorlabs, DMSP680B), all critical properties for AO-OCT performance. To improve usability, the DM was custom cut along an elliptical curve and glued to the conic mount. This streamlined design reduced possible contact with the subject’s face (Fig. 1, see Supplement 1^{(3.3MB, pdf)}, Figs. S4-S6 for design drawings). By angling the DM at 42° instead of the standard 45° relative to the optical axis, a 6° angle was created between the dichroic and the planar mirror that directs the AO-OCT imaging beam to the dichroic before entering the eye. This adjustment shifts the planar mirror farther from the subject, extending the WD between the mirror and the subject’s chin by 13 mm (Fig. 1).

Setting the distance between L2 and L3 to f₂ + f₃ (i.e., 321.4 mm) results in infinite conjugate operation (0 D vergence exiting the subsystem) and a sharp view of the fixation target for emmetropic subjects. Adjusting the L2-L3 distance (which acts as a Badal optometer) provides correction of hyperopic (larger L2-L3 distance) and myopic (smaller L2-L3 distance) refractive errors, the range of which is calculated in the next section. For our Badal design, we ideally wanted to position the back focal plane of the Badal lens L3 (F₃) at the subject’s front nodal point (approximate posterior surface of the eye’s crystalline lens). As shown in Fig. 1 (purple box), the entire fixation-stimulus channel, apart from L3, was mounted on a motorized translation stage (Unislide MA40 series, Velmex, 5” travel range), which allowed this portion of the channel to axially translate towards or away from L3 and therefore adjust the amount of refractive error correction. Although L3 may appear to be orientated incorrectly in our schematic (Fig. 1), we found that this orientation provided a sharper view of the fixation target for the subject.

2.2. Badal optometer design considerations

We wanted our Badal design to ideally position the back focal point of the Badal lens, L3, (Fig. 1, purple box) at the front nodal point of the eye. This arrangement ensures that both the angular field and the angular position of the fixation target on the retina are unaffected by the eye’s axial length provided the image is in focus, which holds true for our application. Additionally, this design provides a larger FOV steering range compared to the more typical Badal optometer design where the back focal point is positioned at the front focal point of the eye (spectacle plane), which causes the Badal lens to be further from the eye and decreases the FOV steering range.

However, locating the eye’s nodal points is difficult. Due to this practical limitation and our need for realigning the eye to the same position, we positioned the L3 focal point slightly anterior to the nodal point at the eye’s entrance pupil, which we could easily identify. Given the close proximity of these two points, the angular field and angular location of the fixation target on the retina should remain insensitive to eye length [42,43]. To confirm, we had subjects with varying refractive errors and axial eye lengths visually compare the angular size of the fixation target (set at 3.75°) to a calibrated 3.75° scan from our AO-OCT system.

We located the subject’s entrance pupil using a separate video camera that live streams an oblique view of the subject’s pupil to a monitor, that was pre-calibrated with a mark indicating the exact depth location of the L3 focal point. During imaging, the center of the subject’s pupil was aligned to the marked position on the monitor as well as the center of the Shack-Hartmann wavefront sensor on the AO computer monitor (see Liu et al. [9] for AO description). The subject’s head position was controlled using a bite bar mounted to an XYZ motorized translation stage system under remote control. Additionally, we found that reflections from the cornea and crystalline lens as seen on our monitor straddled the position of the subject’s pupil and were used to facilitate alignment to reliably and repeatedly place the subject’s pupil at the focal point of L3.

For our system, we wanted to achieve a refractive error correction range of 25 D to accommodate the distribution of spherical refractive errors present in the population, excluding only the most extreme cases [47–49]. To determine the necessary translation range of the Badal system required to correct this range, we used the Newtonian form of the lens equation:

x x^{'} = - {f_{3}}^{2},

(1)

where x is the distance from the front focal point (F₃) to the object (i.e., image formed by L2 at F₂’), x’ is the distance from the back focal point (F₃’) to the image formed by L3, and f₃ is the focal length of the Badal lens L3. Distances x and x’ follow standard sign convention being negative when measured to the left of F₃ and F₃’, respectively (i.e., opposite to the direction of light travel), and positive when measured to the right of F₃ and F₃’, respectively (i.e., in the direction of light travel).

As 1/x’ corresponds to the image wavefront vergence L at F₃’, and similarly the eye’s spherical refractive error, it follows that:

x = - L \cdot f_{3}^{2} .

(2)

This equation shows that the vergence L at the eye’s front nodal point is linearly proportional to the translation distance x. Varying the travel distance of the Badal optometer to correct different refractive errors is illustrated in Fig. 2.

Fig. 2. — **Illustration of how to calculate the stage travel distance x required to correct for different subject refractive errors L.** (top) To bring the fixation target into sharp focus on the retina of an emmetropic eye (L = 0 D), a 0 D vergence wavefront must enter the eye (L7). This is accomplished by axially translating the back focal point of lens L2 (F₂’) to coincide with the front focal point of lens L3 (F₃), causing the optometer image to form at infinity (x’ = ∞). (middle) To bring the fixation target into sharp focus on the retina of a myopic eye (L < 0 D), the image formed by L2 at F₂’ must be shifted past F₃ towards the eye by a distance x (> 0), thereby imparting a negative vergence at the eye. (bottom) To bring the fixation target into sharp focus on the retina of a hyperopic eye (L > 0 D), the image formed by L2 at F₂’ must be shifted to the left of F₃, i.e., farther from the eye, by a distance x (< 0) in order to impart a positive vergence at the eye.

Using a 14 D Volk lens for L3 and targeting a 25 D corrective range, Eq. (2) predicts a necessary translation range of 5” (127 mm), which is achieved by our motorized stage (Fig. 1). Because the distribution of refractive errors in the general population is skewed towards myopia [47–49], we biased the 25 D correction range from −17.5 to +7.5D. Note that in our Badal design, L is the vergence at the eye’s nodal point rather than at the spectacle plane, which is the conventional reference for reporting spherical refractive error. While referencing the spectacle plane is not necessary for our purposes, L can be converted to the spectacle plane vergence by accounting for the distance between the spectacle plane and the eye’s front nodal point. For example, using a 22.27 mm separation distance (as in the Bennett and Rabbetts schematic eye [50]), vergences of −7 D, −3 D, 0 D, + 3 D, and +7 D at the nodal point correspond to −8.3 D, −3.2 D, 0 D, + 2.8 D, and +6.1 D at the spectacle plane, respectively.

2.3. Stimulus subsystem design

The design of the stimulus subsystem (Fig. 1) was guided primarily by the optical requirements of the light stimulus, namely its wavelength, exposure duration, power level, and beam uniformity. Our experiences with an earlier design, previously reported in several studies [17,20,32], further clarified these requirements and revealed areas for improvement to make a more powerful and versatile stimulus subsystem suitable for a wider range of clinical and basic science applications. Based on these considerations and our motivation to accurately characterize functional retinal biomarkers as described in the Introduction, we wanted the new light stimulus to uniformly illuminate a 2° diameter patch of the retina with a short, but adjustable, flash (in milliseconds) of monochromatic light (< 1 nm) at any wavelength across the visible spectrum (400 to 700 nm), and at power levels up to 1 mW entering the eye.

While not all of these specifications were fully achieved, we came close by using a high-power supercontinuum laser combined with an acousto-optic tunable filter (AOTF), ultra narrow bandpass filters in two motorized filter wheels, a mechanical shutter, a multimode optical fiber, and optics designed to efficiently deliver light through the fixation-stimulus channel. Details of these components and their performance in the channel are provided below.

We considered a wide range of light source types, including quartz tungsten halogen and arc lamp tunable light sources, laser diodes, superluminescent diodes, light emitting diodes (LEDs), dye lasers, and supercontinuum lasers. We found quartz tungsten halogen and arc lamp tunable light sources attractive because they are tunable across the full visible spectrum and offer high spectral purity. However, their diffuse emission and extended source size makes them inefficient at coupling light into the eye, resulting in power levels well below that required for our stimulus subsystem. Laser diodes and superluminescent diodes provide the necessary spectral purity and high power, but their wavelength coverage is limited, and they are cumbersome and inefficient to combine. LEDs, which we have previously used [17,20,32], also have limited coverage of the visible spectrum, lack spectral purity (bandwidths >> 2 nm), and are similarly challenging to combine. High-end dye lasers provide the necessary wavelength range, spectral purity, and high power, but they are high maintenance and use dyes that are often toxic. In contrast, supercontinuum lasers offer all the performance advantages of high-end dye lasers without the maintenance and toxicity concerns.

We chose the SuperK Fianinum FIU-15 supercontinuum laser (NKT Photonics) primarily because of its high power output across the full visible wavelength range (350–850 nm), including the highest power level at shorter wavelengths (400–500 nm) among commercially-available supercontinuum lasers. We spectrally filtered and temporally controlled the output of the supercontinuum using an acousto-optic tunable filter (AOTF) (NKT Select, NKT Photonics, VIS 1×, 430-670 nm, 0.5–1.8 nm spectral bandwidth full-width half-maximum (FWHM), > 10 dB sidelobe suppression). This single device allowed us to control the stimulus wavelength, power level, and pulse duration, giving us considerable flexibility in tailoring these parameters for specific experiments but at the expense of a reduced spectral range: 430-670 nm. Although we frequently varied the stimulus wavelength and power level, we typically used a fixed exposure duration of 5 ms, which was brief enough to produce a fast and robust cone ORG response.

Stimulus power was controlled by NKT CONTROL software and modulated using a fast wavelength switching signaling interface (NKT Command) and radio frequency (RF) driver hardware connected to a delay generator (BNC 555 series). This setup allows modulation rates up to 2 kHz (corresponding to exposure durations ≥ 500 µs). Although the NKT COMMAND can support even higher modulation rates, the pulse shape progressively deviates from an ideal flat-top profile as the exposure duration decreases to below 500 µs. For exposures less than 500 µs, the rise and fall times (< 6 μs, measured between 10% and 90% of the steady-state flash intensity) occupy a greater proportion of the pulse duration, producing rounded leading and trailing edges.

While the AOTF controlled the wavelength selection, power modulation, and spectral width (< 2 nm) of the output, the > 10 dB side lobe suppression of the device and the residual white light that leaked from the device could compromise experiments requiring strict stimulus monochromacy. To address this, we introduced two additional devices into the stimulus path that substantially suppressed the sidelobes and blocked the AOTF leakage light. Twelve ultra narrow bandpass filters (1.0 - 2.5 nm spectral bandwidth) from Alluxa and Semrock MaxLine stock were installed in a computer-controlled filter wheel (FW212C, Thorlabs, Ø0.5”, 12 filter positions) following the AOTF (see Fig. 1). Each filter achieves > 60 dB side lobe suppression and eliminates the white light output of the AOTF, except within its narrow bandpass. The 12 bandpass filters spanned most of the visible spectrum with wavelengths at 434.1, 457.5, 472.9, 487.8, 514.4, 543.5, 568.4, 589.0, 606.8, 632.2, 646.3, and 661.2 nm. An exception to the > 60 dB suppression was the 434.1 nm filter, which had > 40 dB suppression.

Note that the 12 wavelengths correspond to the actual wavelengths we empirically determined to maximize transmission through the bandpass filters and do not necessarily represent the center wavelengths specified by the filter manufacturer. Although these bandpass filters were highly effective at suppressing sidelobes of wavelengths immediately adjacent to the specified center wavelength, some filters did not achieve > 60 dB suppression across the entire visible spectrum. Therefore, we cascaded these filters with a second set of bandpass filters (BrightLine, AVR Optics, FF01-458/64, FF01-540/80, and FF01-640/20), which transmitted their center wavelengths with high efficiency (> 90%), while suppressing shorter or longer visible wavelengths that the ultra-narrow bandpass filters could not fully suppress. This second set of filters was installed in a second computer-controlled filter wheel (FW212C, Thorlabs, Ø0.5”, 12 filter positions), placed after the first one (see Fig. 1).

A mechanical shutter (LS6, Vincent Associates) was installed after the filter wheels and opens only during the 5-ms stimulus interval, effectively blocking the small amount of residual leakage light emitted from the AOTF between stimulus flashes. The shutter also provides an independent safety measure in the event of a system failure, such as the AOTF failing in the “on” state, allowing a continuous beam of supercontinuum light to pass through. To properly synchronize these two devices and account for the shutter opening time, the pulse emitted by the AOTF needed to be delayed by 2.09 ms. Our delay generator was triggered only after receiving a relayed trigger signal from our Arduino UNO device sent from our OCT control computer ∼1 s upon initiating AO-OCT video acquisition (50 AO-OCT volumes / video at a 10 Hz volume rate). Upon receiving the Arduino trigger signal, the delay generator immediately triggered our shutter to open (and close after 5 ms) and triggered our AOTF (via the Fast Wavelength Switching Signaling Interface (NKT Command)).

After passing through the shutter, the stimulus light coupled into a multi-mode patch cable (FG365UEC-CUSTOM, Thorlabs, core size = 365 µm, numerical aperture (NA) = 0.22, length = 4 m, see section below for fiber selection criteria) using a high NA achromatic lens (F950SMA-A, Thorlabs), which dispersed the light across the thousands of fiber modes and distributed it uniformly over the exit facet of the fiber core. The 365 µm fiber core (retinal conjugate plane) was then imaged onto a 2° patch of retina (∼600 µm diameter for an emmetropic eye) using three pairs of relay telescope lenses (L4-L5, L6-L2, and L3-L7), with the eye optics, L7, serving as the last lens. The stimulus beam aligned to the fixation subsystem axis.

Anticipating losses in stimulus power from combining the stimulus subsystem with the fixation subsystem and AO-OCT system, we used achromatic doublets with visible-light AR coatings, a folding planar mirror with dielectric coating tuned to visible light, and a 10:90 (T:R) beamsplitter to maximize stimulus power throughput. Light exiting the multi-mode patch cable was collimated using a microscope objective L4 (DIN 20×, Newport, NA = 0.4, f = 8.55 mm) and beam expanded by two achromatic doublets L5 (AC254-030-A, Thorlabs, f = 30 mm, Ø1”) and L6 (PAC055, Newport, f = 125 mm, Ø1”) spaced 155 mm (f₅ + f₆ distance) apart. The collimated beam exiting L6 was redirected via an elliptical dielectric mirror (BBE1-E02, Thorlabs) to our beamsplitter (BSX16, Thorlabs, Ø2” 90:10 (R:T) UVFS, 400-700 nm, thickness = 8 mm) that coupled our stimulus and fixation subsystems together. The reflected light from the 10:90 beamsplitter passed through the remaining relay telescope consisting of L2 and L3 before reaching the eye.

2.4. Considerations in selecting the multi-mode fiber and lenses in the stimulus subsystem

The specific design of the stimulus subsystem is constrained by how it optically integrates with both the subject’s eye and the fixation subsystem. In particular, because the stimulus light must pass through the Badal system, it needs to be introduced upstream of lens L2 in the fixation subsystem, but downstream of L1, due to inherent size mismatches between the fiber tip and the microdisplay. To accommodate this, we used three lenses in the non-common part of the stimulus subsystem, in contrast to the single lens used in the non-common part of the fixation subsystem (see Fig. 1). In this configuration, the first two lenses of our stimulus subsystem form a relay telescope, while the third lens, together with L2 (from the fixation subsystem) form a second telescope. This arrangement provides greater flexibility in meeting the key design requirements of the stimulus subsystem:

1.
D_fiber × M_{stim_chan} = 600 μm, where D_fiber is the fiber core diameter and M_{stim_chan} is the linear magnification of the stimulus subsystem. M_{stim_chan} = (f₅/f₄) × (f₂/f₆) × (f₇/f₃), with f₂ and f₃ specified by the requirements of the fixation subsystem.
2.
Stimulus beam diameter D_beam < D_eye (pupil diameter) to avoid vignetting.
3.
LI_fiber < LI_eye = 60 μm·rad, where LI_fiber and LI_eye are the Lagrange invariant of the multi-mode fiber and the eye.
4.
NA_fiber < NA_L4, where NA_fiber and NA_L4 are the numerical apertures of the multi-mode fiber and the first lens (L4) collecting the fiber light.
5.
Distance between L6 and L2 (f₆ + f₂) is sufficiently long to redirect the stimulus beam into the fixation subsystem via a folding mirror and beamsplitter.
6.
Length and size of the stimulus subsystem are compact enough for mounting onto the motorized Badal system.
7.
Design is limited to available stock lenses, mirrors, beamsplitters all ≤ 2” diameter, and stock multi-mode fibers.

The first and second requirement (1 & 2) specifies that the fiber tip (core) illuminates a 2° retinal patch while maintaining a beam diameter at the pupil (D_beam) less than the pupil diameter (D_eye = 6.7 mm) to avoid vignetting. This is achieved by adjusting four free parameters: D_fiber, f₄, f₅ and f₆. The third requirement (3) considers the optical Lagrange invariant (LI), defined as the product of ynu, where y is the object or image height, n is the refractive index and u is the ray angle. The product nu also corresponds to the numerical aperture (NA) when u is small. The Lagrange invariant remains constant through an ideal, lossless optical system. For a fixed Lagrange invariant of the eye (LI_eye), a larger Lagrange invariant of the stimulus subsystem’s fiber (LI_fiber) results in clipping or vignetting at the pupil plane or overfilling at the retina plane, both leading to light loss. Thus, to minimize this loss, the fiber’s Lagrange invariant (LI_fiber) was chosen to be less than the eye’s Lagrange invariant (LI_eye), while also conforming to requirements 1 & 2. The LI of the eye is calculated to be LI_eye = y’n’u’ = 60 µm·rad, where y’ = 300 µm (half the 2° diameter stimulus patch), n’ = 1.33, u’ = D_eye/2/F’_eye, D_eye = 6.7 mm (pupil diameter), F’_eye = 22.22 mm (eye effective focal length). Therefore LI_fiber = D_fiber × NA_fiber < 60 μm·rad gives us a second equation with two free parameters. While the five free parameters (D_fiber, NA_fiber, f₄, f₅ and f₆) give us considerable flexibility, they are further constrained by the requirements for the first lens (NA_fiber < NA_L4) (4), the long distance between L6 and L2 to combine the stimulus and fixation subsystems (5), the compact size of the stimulus subsystem for mounting on the movable arm of the Badal system (6), and the limited commercially-available fibers and lenses (7).

Through iteration, we converged on a microscope objective (20×, NA = 0.4, f = 8.55 mm) for L4 and two 1” achromatic lenses for L5 (f = 30 mm) and L6 (f = 125 mm), which when combined with L2, L3 and L7 (the eye) provided a total magnification (M_{stim_chan}) from fiber tip to retina of 1.64. Under these conditions, a fiber with D_fiber = 365 μm produces a 2° stimulus patch on the retina.

If a larger stimulus field is required, two general strategies are possible. The first is to increase the linear magnification of the stimulus subsystem by adjusting the focal lengths of lenses L4-L6. For example, reducing the focal length of objective lenses L4 or L6, or increasing that of L5 can increase the FOV. This can be realized by replacing the lenses or, for greater flexibility, by incorporating tunable lenses. The second strategy is to increase the fiber’s diameter (D_fiber) using a fiber with a larger core but similar NA, although suitable options may be limited. To properly balance design requirements for a larger stimulus field while still optimizing light throughput, refer to the relevant design criteria outlined in Requirements 1-7 above.

The multimode fiber we selected (FG365UEC-CUSTOM, Thorlabs, core size = 365 µm, numerical aperture (NA) = 0.22, length = 4 m) has a LI = 40.1 µm·rad, which is below the 60 µm·rad upper limit that we specified for the eye and results in a beam diameter at the eye pupil of 4.6 mm. We did not choose a smaller fiber NA, as doing so would further reduce the beam size at the eye pupil, making the stimulation of photoreceptors potentially more sensitive to eye alignment (i.e., lateral displacement of the stimulus beam) due to the Stiles-Crawford effect [51]. Moreover, although fibers with numerical apertures up to 0.33 would still satisfy the LI constraint for the eye, such fibers are less readily available and more prone to partial occlusion of the stimulus beam due to eye motion.

2.5. System validation, AO-OCT imaging, and optoretinography stimulation protocols

The use of two proprietary commercial lenses (Volk Optical lens and 7-element Canon camera lens) in the fixation-stimulus channel prevented us from evaluating the theoretical performance of the channel using commercial optical ray tracing software. Nevertheless, we experimentally measured key performance parameters of both the fixation and stimulus subsystems to validate their performances. For the fixation subsystem, we determined the visual field steering range of the fixation target, assessed image quality and distortion across the steering range, confirmed the physical working distance between the fixation-stimulus channel and the subject’s eye, and assessed target brightness. For the stimulus subsystem, we assessed its ability to produce (1) spatially flat and temporally top-hat intensity profiles, (2) spectrally selective (narrowband) light at 12 wavelengths across the visible spectrum, and (3) short high-power flashes. By evaluating these key performance parameters, we confirmed that the system met the prescribed design requirements of an integrated fixation-stimulus channel for AO ophthalmic use.

To determine the FOV of our fixation subsystem, we first calibrated it using a previously calibrated 0.8°×3.75° rectangular raster pattern created by the AO-OCT imaging beam. The subject was instructed to adjust the fixation target position until it was centered within this raster. To calibrate the retinal pixel-per-degree scale, the subject then adjusted the fixation target height until it matched the slow-axis extent of the scan (3.75°), allowing us to determine the number of pixels corresponding to 3.75 degrees of retinal eccentricity. Using a relatively large scan angle (3.75°) improved accuracy of the task, while keeping the target sufficiently small to minimize distortion effects within the fixation subsystem (discussed below).

Next, to map the limits of the fixation subsystem’s FOV, the subject moved the fixation target toward the edge of their visual field until the center of the target was no longer visible. The coordinates at which visibility was lost were recorded at four positions—two along the horizontal meridian and two along the vertical. To test that the FOV of our fixation design was insensitive to axial length of the eye, we repeated this calibration procedure in three subjects with varying refractive errors and axial lengths: S1 (ID199): −0.5 D spherical refractive error, 24.14 mm length; S2 (ID154): −3 D, 25.24 mm; S3 (ID321): −5.75 D, 26.79 mm). Note that a fourth subject, S4 (ID204), participated only in the functional imaging experiments described later. Subject ID numbers refer to our internal subject identification system.

Image sharpness and distortion of the fixation target were assessed using a white-on-black grid pattern generated by the microdisplay and imaged through the fixation subsystem with a digital color camera (EOS R10 Mirrorless Camera with 18-150 mm Lens, Canon) serving as a substitute for the subject’s eye. A 20 × 15 box grid (twenty-one vertical and sixteen horizontal lines spaced 40 pixels apart) was displayed on the microdisplay. Using the pixel-per-degree calibration described above, the FOV observed by our subjects was consistent with the horizontal and vertical FOV of the image captured by our digital color camera. Image sharpness was evaluated by visually inspecting the lines for transverse chromatic aberrations and overall blurring with field angle. Distortion was measured along the horizontal and vertical meridians using the width of the central square (near white circle in Fig. 3) as a reference, assuming it is undistorted. This central reference was then used to extrapolate the expected “distortion-free” grid line positions out to the edge of the FOV. The locations of these ideal lines relative to the center of the image (Ideal Line Distance) were then compared to the actual line locations in the acquired image relative to the center of the image (Actual Line Distance). A distortion percentage was calculated using

Distortion % = 100 \times (Ideal Line Distance - Actual Line Distance) / Ideal Line Distance .

(3)

Fig. 3. — Characterizing the fixation subsystem. (A) View of the microdisplay with grid pattern, as captured by a color digital camera aligned with the fixation subsystem and serving as a substitute for the subject’s eye. Grid pattern is observed over an oval 39.0° (H) × 36.6° (V) FOV. (B) Distortion in the grid pattern is measured as a percentage, Eq. (3), and plotted as a function of retinal eccentricity across the horizontal and vertical meridians as denoted by the cyan and yellow dashed lines in (A).

Next, we evaluated the spatial, temporal, and spectral performance of the stimulus subsystem. A flat stimulus beam profile assures uniform illumination of photoreceptors within the designed 2° retinal patch. To quantify flatness, we acquired images of the beam’s cross section at the retinal plane (conjugate to the multimode fiber tip) using a model eye composed of an achromatic lens (AC254-30, Thorlabs, f = 30 mm) and a digital monochrome camera (DCC1545 M, Thorlabs) inserted at the location of the eye. We then measured the coefficient of variance (COV) from these images.

A top-hat profile for the stimulus flash with fast rising and falling edges assures a linear relationship between stimulus exposure duration and total photons incident on the retina. To quantify the flash profile, we recorded the stimulus trace for our double flash ORG protocol (see details below and in Supplement 1^{(3.3MB, pdf)}) using a photodetector (New Focus, Inc., 2001) connected to a digital oscilloscope (Tektronix, DPO 4034B). Measurements of the top-hat trace included rise and fall times, and width. We also measured the throughput efficiency, from AOTF to subject eye, and the maximum power level achievable incident at the cornea using a Newport 843-R, 818-ST2/DB power meter (±1% calibration uncertainty for visible wavelengths). These measurements were repeated at each of the 12 wavelengths from 1-100% RF in increments of 1% RF.

To assess spectral performance, we confirmed the center wavelength on the eye for each of the 12 stimulus wavelengths using a handheld spectrometer (StellarNet Inc.). We also quantified the effectiveness of the ultra-narrow bandpass filter transmission spectra to suppress the AOTF leakage light. This was estimated by applying the ultra-narrow bandpass filter transmission spectra, obtained from the manufacturer, to the AOTF background spectra measured using the spectrometer.

Finally, we performed two imaging experiments to investigate cellular-level structure and function using the Indiana AO-OCT system. One targeted the inner retina, and the other the outer retina. Both were performed at 19.5° retinal eccentricity – the outermost extent of the fixation subsystem’s FOV. This allowed us to evaluate system performance at peripheral retinal locations we had not previously imaged. Imaging was performed on the right eyes of Subjects S2 and S4. Prior to imaging, one drop of 1% Tropicamide was administered to dilate the subject’s pupil.

Briefly, the Indiana AO-OCT system operated at A-scan rates of up to 1 MHz [41,52], achieving lateral and axial resolutions of 2.4 µm and 4.7 µm, respectively. The system’s AO loop ran at rates up to 233 Hz with a loop gain of up to 1 [9,53]. Our global registration algorithm [54] provided extended FOV images by combining images that were displaced by retinal motion.

For structural imaging, we acquired images of retinal ganglion cells and other inner retinal cellular-level structures followed by imaging of photoreceptors and retinal pigment epithelium (RPE) cells in the outer retina. For functional imaging, we performed optoretinography on cone photoreceptors using a new double-flash stimulus protocol, consisting of a 434.1 nm flash (151 µW at the cornea, 1.6% V(λ)-weighted bleach) followed one second later by a 606.8 nm flash (743 µW at the cornea, 43.0% V(λ)-weighted bleach). The ORG responses were used to classify the cone spectral types (S, M, and L). All visible light stimulation was delivered at exposures at least 12× below the ANSI safety limit after accounting for the additional AO-OCT imaging beam [55].

Additional details of these experiments and how stimulus bleach fractions were calculated can be found in the Supplement 1^{(3.3MB, pdf)}.

3. Results

We report key performance metrics of our fixation-stimulus channel, along with demonstrations in subjects using the Indiana AO-OCT system integrated with this channel.

The system’s working distances (WDs) measured from the tip and center of the dichroic mirror (DM) to the cornea of one subject were 27 mm and 42 mm, respectively. These distances have proven sufficient to accommodate over 50 subjects imaged to date, including individuals with deep-set eyes and mild facial hair.

The fixation subsystem achieved a 39.0° × 36.6° FOV centered on the fovea across the three subjects (S1, S2, and S3), consistent with measurements obtained using a digital color camera as a model eye (Fig. 3(A)). FOV calibration in all three subjects yielded the same angular resolution of 16.4 pixels/degree on the fixation display, as expected based on the Badal principle, which predicts that angular FOV is insensitive to axial eye length.

As acquired by the digital color camera, some color separation of the grid lines is visible in the peripheral regions of the target color image in Fig. 3(A), indicating the presence of transverse chromatic aberrations (TCA) in the fixation subsystem. The color separation matches that reported by the three subjects when viewing the target, suggesting that the TCA originates from the fixation subsystem itself rather than from the color digital camera.

After converting the RGB color image to grayscale and measuring grid spacing, we quantified fixation target distortion as a function of retinal eccentricity (Fig. 3(B)), using the percentage distortion defined in Eq. (3). A positive distortion percentage indicates a shortening of grid line spacing relative to the ideal spacing (barrel distortion), whereas a negative value indicates a lengthening (pincushion distortion).

The average absolute distortion across both meridians was 0.74% ± 0.52% (mean ± SD), with a maximum of 1.9% occurring at the temporal edge (-19.1° horizontal). These distortion levels correspond to absolute angular errors at the retina – computed as (% distortion × retinal eccentricity) – of just 0.07° on average and 0.4° maximum.

Next, we assessed the performance of the stimulus subsystem. We started with quality and flatness of the stimulus beam cross section across the 2° stimulus field (Fig. 4(A)). Our initial measurements revealed a Gaussian beam profile, which we attributed to exciting primarily low-order modes of the multimode fiber. To improve mode mixing and achieve a more uniform distribution of low- and high-order modes, we found that bending the fiber at three adjacent points near its entrance effectively produced a uniform beam intensity at the retinal plane of our model eye. The COV of the elevated (top-hat) portions of the horizontal and vertical intensity profiles of the beam was less than 5%, indicating good uniformity. We found the mechanical bending of the fiber to be highly stable, with no detectable changes in the spectral output or beam profile of the stimulus light over time. Although small fluctuations in optical power (< 5%) were observed, similar variability in measurements taken before and after coupling into the multimode fiber indicates that this instability likely originated within the supercontinuum source rather than the fiber itself.

Fig. 4. — Characterizing the stimulus subsystem. (A) Image of a 606.8 nm stimulus beam at the retinal conjugate plane of a model eye. Horizontal and vertical intensity profiles reveal a quasi-flat beam profile with a COV of < 5%. (B) AOTF background light before (black) and after filtering with the 606.8 nm ultra-narrow bandpass filter (orange). The filter suppresses AOTF leakage light by > 60 dB, achieving a high degree of pure narrowband light. (C) Double-flash ORG imaging protocol showing the electrical signal trace from the Arduino (black) that triggers a delay generator that triggers both the AOTF and a mechanical shutter to emit two high intensity 5 ms flashes spaced 1 s apart (red). Further details of the ORG protocol are provided in Methods and the Supplement 1^{(3.3MB, pdf)}. (Inset) A magnified view of a single 5 ms stimulus flash reveals a pristine top-hat intensity profile. (D) Power (µW) measured at the cornea for all 12 wavelengths as a function of AOTF driver RF power (%). Colored lines denote different wavelengths in the visible spectrum, though the line color may not match the actual color of the wavelength. (E) Shows 12 test wavelengths in (D) sampled across the visible spectrum and overlaid on Stiles & Burch-based 10° human L (red), M (green), and S (blue) cone (pre-receptoral) quantal sensitivity (S_λ) curves [56–58]. Colored dashed lines nominally match the wavelength color.

Figure 4(B) demonstrates the effectiveness of ultra narrow bandpass filtration in reducing AOTF leakage light and spectral sidelobes — secondary energy peaks at wavelengths adjacent to the filter’s center wavelength — by more than 60 dB. In Fig. 4(C), we show our trigger signal (black) and the measured 434.1 nm and 606.8 nm 5 ms flashes (red), separated by 1 s. Timing uncertainty of the stimulus flash onset was ±0.6 ms (±SD) relative to the start of the AO-OCT volume acquisition. The inset highlights a zoomed-in view of the recorded stimulus flash, showing that we achieved a 5 ms flash. Rise and fall times of the flashes were < 6 μs (10% to 90% of the steady value of the flash), producing a pristine top-hat profile. Finally, we achieved stimulus power levels ranging from 0 to 220 µW at 434.1 nm wavelength and 0 to 1,210 µW at 646.3 nm wavelength (Fig. 4(D)). The dashed-colored lines in Fig. 4(E) show our 12 test wavelengths (spectral samples) across the visible spectrum relative to the 10° quantal and pre-receptoral (i.e., removes lens and macular pigment contribution) cone fundamentals based on Stiles & Burch’s color matching functions (CMF) [56–58].

Table 1 summarizes the key parameters of the integrated fixation-stimulus channel.

Table 1. Summary of fixation-stimulus channel parameters.

Fixation Subsystem Specifications
Field of view	39.0° (H) × 36.6° (V)
Working distance (from DM center to cornea)	42 mm
Badal optometer correction range	−15 D to +7.5 D (25 D total)
Travel range	5” (127 mm)
Display FOV calibration	16.4 pixels / degree
Contrast ratio	10,000:1
Distortion error	0.07° on average and 0.4° maximum
Stimulus Subsystem Specifications
Exposure duration	> 0.5 ms (5 ms, typical)
Retinal stimulus size	2° dia.
Beam diameter (at cornea)	4.6 mm
Beam profile flatness at retina	COV < 5%
Spectral tuning range	430–670 nm
Monochromacy	< 2 nm FWHM
Sidelobe suppression	> 60 dB except > 40 dB @ 434. 1 nm
Maximum stimulus power	0.22 mW @ 434.1 nm and 1.21 mW @ 646.3 nm

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As shown in Fig. 5(A), the steering FOV of the fixation subsystem (red oval) is overlaid on a wide-field 55° clinical SLO image and extends beyond the optic disc and the vascular arcades (>18° retinal eccentricity). An OCT angiography (OCTA) image is also overlaid to assist in localizing vasculature visible in our AO-OCT images (Fig. 5(B)–(D)) and used to identify the imaged location. The corresponding OCT B-scan provided a foveal reference (identified by the hyper reflective spot where the retina was minimally thin) and used to calculate retinal eccentricity. The small offset between our measured retinal eccentricity (19.5°) and that of the SLO image (∼18.4°, yellow box) is discussed in the Discussion section. Fig. 5(B)–(D) show en face views of the outer retina of photoreceptor outer segments, rod outer segment tips, and retinal pigment epithelium (RPE) cells at the 19.5° edge of the fixation subsystem’s FOV. After image registration and averaging, the FOV of the final AO-OCT images increased from 0.8° × 1° to 1.2° × 1.2°.

In the same subject and at the same retinal eccentricity, Fig. 6 shows en face projections of the inner retina showing retinal ganglion cells, displaced amacrine cells, retinal vasculature, and nerve fiber bundles. Following image registration and averaging, the horizontal FOV of the final AO-OCT images increased from 1.4° to 1.59°, while the vertical FOV remained unchanged. Note that in Fig. 6 the images were slightly cropped in the vertical dimension from 1.5° to 1.45°.

Lastly, we used the functional response of cone photoreceptors (the cone ORG) to a double-flash of 434.1 nm and 606.8 nm stimuli light to classify cones by their spectral type (S, M, or L). As shown in Fig. 7, measurements were made in subject S4 at 19.5° retinal eccentricity. Due to the short cone outer segment length (8–13 µm) (Fig. 7(B), top), we adapted our previously reported processing of optical path length changes (ΔOPL) [17,32]. By isolating the inner segment / outer segment junction (IS/OS) and cone outer segment tip (COST) reflections of individual cones and computing a ± 2 axial pixel complex average at each individual cone’s IS/OS and COST reflection, ΔOPL was calculated for the shortened outer segments. The ±2 axial pixel range nominally spanned the axial resolution of the AO-OCT system (4.7 µm in retinal tissue). Classified S, M, and L cone responses are shown in Fig. 7(A) as blue, green, and red traces, respectively, obtained using principal component analysis and a Gaussian mixture model as described elsewhere [17,32]. Figure 7(B), bottom, shows the S, M, and L cone classification results overlaid on the en face projection of the cone outer segments, revealing the trichromatic cone mosaic. The measured L:M:S cone ratio was 52%:38%:10%. The AO-OCT B-scan highlighted in Fig. 7(B) also revealed cones exhibiting multimodal IS/OS and COST reflections, likely arising from the waveguiding properties of cones and their larger apertures at this high retinal eccentricity [59].

Fig. 7. — **AO-OCT optoretinography and cone classification in subject S4 at 19.5° temporal to the fovea.** (A) ORG responses from 255 cones are shown following a double-flash stimulus (434.1 nm and 606.8 nm). The ORG captures optical path length changes (ΔOPL) within the outer segment, which are used to classify cones into S, M, or L spectral types, color-coded as blue, green, and red, respectively. (B) An AO-OCT B-scan image shows several cone cells with short outer segments (∼12 µm) and enlarged IS/OS and COST apertures, as expected at this far retinal eccentricity. The cone classification results are mapped onto an *en face* projection of the cone outer segments, revealing the trichromatic cone mosaic. Inward pointing arrows indicate the location of the B-scan image (top). Abbreviations: IS/OS, inner segment / outer segment junction; COST, cone outer segment tip; ROST, rod outer segment tip; RPE, retinal pigment epithelium; BM, Bruch’s membrane. Scale bars are 50 µm.

4. Discussion

We developed an integrated fixation-stimulus channel focused on key performance attributes to enhance the utility of AO ophthalmoscopy. Here, we discuss our findings on the main features of the fixation and stimulus subsystems, the key trade-offs made to meet design requirements, differences over previous designs [23], current limitations of the channel, and examples demonstrating its usefulness within the Indiana AO-OCT system.

4.1. Fixation subsystem performance and key design trade-offs

For the fixation subsystem, we converged on a cascade of two lens-based relay telescopes, with the final telescope in a Badal configuration with the eye. All lenses were stock components, with no lens exceeding 52 mm (∼2”) in diameter. As expected, the Badal design provided an angular FOV that was insensitive to the subject’s refractive error and axial length. We achieved a fixation target FOV of ±19.5° × ±18.3°, enabling imaging beyond the optic disc and about 1.5 × further out in the peripheral retina compared to the previous fixation subsystem in the Indiana AO-OCT system [15].

The accuracy of our fixation target placement exceeded typical eye fixation stability reported in the literature. System-induced distortion of the fixation channel produced angular errors at the retina with an average absolute error of 0.07° and a maximum error of 0.4° across both horizontal and vertical meridians (see Results). These values are well below the fixation variability observed in clinical populations, where fixation points typically fall within 0.50° and 0.87° of the target location for the 63% and 95% confidence intervals, respectively [60]. Thus, the systematic fixation error induced from channel distortion are small compared to the intrinsic random fixation error observed in these clinical populations. Moreover, the offset between our predicted retinal location (19.5°) and that seen in the Spectralis SLO image (Fig. 5(A)) is consistent with mapping distortions reported by Huang et al. [26], who found errors up to ∼0.8° at 10° eccentricity between the Spectralis and their AO-SLO system. Extrapolating this relationship predicts errors exceeding 1° at ∼20° eccentricity, consistent with the small offset observed here. For applications requiring higher spatial accuracy, these fixed errors can be corrected by adjusting the fixation target position on the microdisplay at each eccentricity.

We did not directly measure the luminance of the fixation target at the eye, however, the brightness required for comfortable viewing was empirically found to be 67% of the microdisplay’s maximum output. This indicates that our design provides ample brightness reserve, even with the use of a 10:90 beamsplitter, which reduced the display brightness by ∼90%.

With as little as 5 inches of motorized translation of the Badal configuration, we achieved a 25 D correction range, which we biased from −17.5 D to +7.5 D to accommodate the myopic skew of the general population [47–49]. While the AO correction range of the Indiana AO-OCT system is more limited (currently ±6.5 D), this extended Badal range will benefit AO ophthalmoscopes with larger correction capabilities, which will be particularly valuable for imaging highly myopic eyes. Additionally, we found the wide dioptric range helps to simplify setup and alignment of the fixation subsystem within the AO ophthalmoscope, benefits with essentially no added cost.

One of our most important design considerations was balancing the trade-off between subject working distance, WD, and the fixation subsystem’s FOV. As illustrated in Fig. 8, the working distance is proportional to the focal length of the (Badal) lens positioned immediately in front of the eye (L3), whereas the FOV of the microdisplay at the retina is inversely proportional to that focal length. Thus, increasing the working distance by adjusting this focal length means reducing the FOV. FOV is also proportional to the diameter of L3 (assuming it is the limiting diameter in the subsystem), thus a larger lens L3 provides a larger FOV.

Fig. 8. — **Fixation subsystem tradeoff between subject working distance and field-of-view.** (Top) Schematic shows how a shorter focal length Badal lens L3 decreases the working distance (WD) between the dichroic mirror (DM) and the eye, while increasing the field-of-view (FOV) on the retina (i.e., farther retinal eccentric locations) for a microdisplay of fixed size. (Bottom) Schematic showing how a longer focal length Badal lens L3* increases the WD, while decreasing the FOV on the retina for an microdisplay of fixed size, reducing the range of retinal eccentric locations to those closer to the fovea. DM, dichroic mirror; F_i, the front focal point of lens Li; F_i’, the back focal point of lens Li. i indicates index of lens.

For practical reasons, we chose to keep the diameter of all optics under ∼2 inches. With this constraint, we found the Volk lens (f_eff = 71.4 mm, entrance pupil diameter = 52 mm) offered a good compromise between working distance and FOV, allowing us to achieve a fixation steering FOV of 39.0° (H) × 36.6° (V) (Figs. 3 & 5) and a working distance of 42 mm from the subject’s eye to the center of the dichroic mirror (Fig. 1). This FOV is approximately 50% wider compared to our previous fixation subsystem in the Indiana AO-OCT system (increase from 13° [15] to 19.5°). The 3D printed conic mount—housing the Volk lens and dichroic mirror—further reduced the risk of contact with the subject’s face and did not vignette any part of the fixation subsystem FOV (see Supplement 1^{(3.3MB, pdf)}, Figs. S4-S6 for design schematics). To date, this arrangement has been sufficient for all subjects we have imaged (n > 50), including those with deep-set eyes and facial hair.

Achieving a significantly larger FOV than that provided by our design for the same working distance will likely require custom optics with diameters greater than 2 inches. Interestingly, as already mentioned in the Introduction, larger FOVs (up to 60° horizontally) have been achieved with an alternative fixation subsystem approach based on a pair of large-diameter spherical mirrors and a pair of orthogonal galvo-tracking mirrors positioned in front of the subject’s eye [25]. Although this design configuration is more complex and less compatible with existing AO ophthalmoscopes, its design is well suited for peripheral imaging.

4.2. Fixation subsystem comparison to literature

It is difficult to compare the performance of our fixation subsystem with that reported by Steven et al. [23]–the only publication we are aware of that is dedicated to this topic–due to fundamental differences in design strategy and intended application. Their design also incorporates a separate fundus camera subsystem while in ours we include a separate stimulus subsystem. Nevertheless, a few comparisons can be made to help elucidate the strengths and weaknesses of each.

Both fixation subsystems share a three-lens optical design using a Badal configuration. However, two of the three lenses differ significantly due to differing design strategies. For their strategy, Steven et al. employed commercial ray tracing software to assess each lens, conducting a rigorous analysis of performance and tolerancing, setting a goal of diffraction-limited performance (including correcting the average longitudinal chromatic aberration and spherical aberration in the eye), and customizing the objective lens positioned directly in front of the fixation target.

In contrast we chose a more simplistic and lower cost approach. For our design, we constrained it to all stock optics with none larger than ∼2” in diameter. However, we considered low-cost high-performance commercial optics, including proprietary ones, which Steven et al. did not consider. Our design was also guided by the specific vision task required of the subject: fixate on a visible target. Because of this, we did not attempt to correct the longitudinal chromatic aberration (LCA) of the eye as it would unlikely improve the subject’s ability to fixate and in fact there is increasing evidence that the eye uses LCA to accommodate [61], which would be important in subjects that cannot be cyclopleged [9]. Similarly, because distortion and TCA are the primary aberrations that directly affect the positional accuracy of the fixation target, we targeted these aberrations in our analysis and evaluated their magnitude relative to the typical fixational accuracy of subjects reported in the literature.

From these complementary design strategies, the three lenses used in the two designs were: (L1) our 7-element Canon camera lens vs their 4-element custom lens, (L2) our paired stock achromats vs their paired stock achromats, and (L3) our aspheric Volk singlet vs their paired stock achromats. Of these, the choice of the Volk lens for L3, instead of paired stock achromats, likely has the greatest impact on overall channel design. The Volk’s shorter focal length (71.4 mm compared to 110.3 mm (35% shorter) was instrumental in allowing us to use ∼2 inch diameter optics (instead of ∼3 inch optics required by Steven et al.). Also, the singlet’s narrower width combined with our custom conic mount resulted in a comparable working distance (42 mm vs ∼45 mm). Note that the ∼45 mm for Steven et al. is our estimate measured from the eye to the dichroic mirror or one half of their reported eye relief, 90 mm, since roughly one half of the relief is likely expended on the 45° dichroic mirror in front of the eye (compare their Fig. 1 to our Fig. 1). Finally, in both designs, L3 serves as the Badal lens. Due to the shorter focal length of the Volk lens, the travel distance required to correct the same refractive error range is reduced by a factor of 2.4×, specifically (110.3 mm/71.4 mm)², according to Eq. (2). Consequently, our system requires only 5” of travel to correct a 25 D range, compared to 12” in the referenced design, resulting in a more compact and cost-effective setup. When configured for an extreme case of a + 7.5 D eye, the total unfolded length of the optical path—from the fixation target to the eye—is approximately 30% shorter in our system (649 mm vs. ∼929 mm), further underscoring the efficiency and compactness of our design.

Note that our design benefitted from the approximately 2 × larger size (either dimension) of the OLED microdisplay compared to the Digital Light Projector (DLP) used by Steven et al. (12 × 9 mm vs 6.9 × 3.9 mm). The larger size reduced the required magnification between the fixation target and the retina, thereby allowing for a longer focal length objective lens in front of the fixation target (our 50 mm vs their 17.9 mm). Overall, our design provided a larger fixation steering FOV: 39.0° (H) × 36.6° (V) compared to 32.2° (H) × 32.2° (V), an improvement of 21.2% (H) and 13.7% (V). This difference enabled imaging at and beyond the optic disc and the vascular arcades.

As already mentioned, the design advantages we gained by using high-performance commercial optics came at the cost of being unable to predict their individual and combined performance in the fixation and stimulus subsystems using commercial ray tracing and tolerancing software. In contrast, Steven et al. were able to leverage these powerful tools to analyze and predict the performance of their design.

Note that neither our design nor Steven et al. corrects for cylindrical (astigmatic) errors, despite the relatively high prevalence of cylindrical error in the population (36.2% for cylinder ≥ 1.0 D) [62]. In our system, we have not observed any noticeable difference in fixation stability among subjects with uncorrected cylinder (some up to 2.5 D cyl). However, in populations where cylindrical correction is necessary, variable cylindrical power devices, such as the Jackson-cylinder [63], could be integrated into the fixation subsystem.

4.3. Stimulus subsystem performance and key tradeoffs

The primary objective of the stimulus subsystem was to deliver intense, uniformly illuminated, short-duration flashes of monochromatic light across the visible spectrum to the retina patch imaged with AO-OCT (∼1.28° diameter = √(0.8² + 1²)). We achieved much of this using a flexible platform centered around the NKT SuperK Fianinum FIU-15 supercontinuum laser with SELECT module. This platform provided millisecond-level control of the flash duration at wavelengths ranging from 430 nm to 670 nm, with maximum power levels at the eye ranging from 220 to 1,210 μW depending on wavelength (Fig. 4(D)). These power levels, delivered in 5 ms flashes, were sufficient to induce several hundred nanometers of change in the photoreceptor ORG, even at large retinal eccentricities where photoreceptor outer segments are short and ORG amplitudes are reduced (Fig. 7). In addition, the system produced temporally precise, top-hat-shaped pulses with sharp rise and fall times (< 6 μs, Fig. 4(C)) and a spatially uniform intensity profile at the fiber tip (conjugate to the retina), with a COV better than 5%. This level of performance is more than sufficient for studies of photoreceptor function. As an ongoing project, we are leveraging the versatility of this stimulus subsystem to deliver high-intensity, near-monochromatic light for characterizing cone spectral sensitivity in the living human eye [35].

Despite these strengths, the stimulus subsystem has several limitations. First, its wavelength range (430-670 nm) does not span the full visible spectrum (400-700 nm), notably excluding wavelengths below 430 nm, which include the peak spectral sensitivity of S cones (Fig. 4(E)). Although wavelengths longer than 670 nm can be accessed using a different acousto-optic tunable filter—such as the nIR1 module (640-1100 nm)—this output is only available from a different port on the SELECT module than the VIS (1x) port used in the current setup.

Another limitation is the SELECT module’s sidelobe suppression, which is limited to > 10 dB. To meet the requirements of our applications, we inserted ultra-narrow bandpass spectral filters to further reduce the sidelobes. While these filters were effective, they restricted the wavelengths we could use and increased cost. For applications not requiring such stringent sidelobe suppression, the full granular wavelength control of the SELECT is accessible.

Our design also employed a multimode optical fiber to achieve a highly uniform intensity profile, but this approach is not readily amenable to generating other illumination patterns or structures, as for example proposed in DLP-based systems [23]. Lastly, the supercontinuum source, along with its supporting hardware and software, add significant cost to the overall system.

4.4. Structural and functional imaging of the peripheral human retina

As illustrative examples, we used our fixation-stimulus channel to perform structural and functional AO-OCT imaging of the human retina at eccentricities extending beyond the optic disc and the vascular arcades (>18° retinal eccentricity). We found that many of our structural and functional tools that we originally developed for retinal locations near the fovea also performed effectively at these larger retinal eccentricities.

At 19.5° temporal eccentricity, structural AO-OCT imaging (Figs. 5 and 7(B)) revealed cone and rod photoreceptors, as well as RPE cells. The cone photoreceptors coarsely sampled the retina, exhibiting enlarged inner and outer segment diameters, and shortened outer segments. These were consistent with histology [64,65]. Interestingly, we observed a cone-to-RPE cell ratio of 1:1 at 19.5°, which to our knowledge, has not been previously reported at such a high retinal eccentricity in the living human retina. This ratio is significantly lower than previously reported in vivo values at more central locations: 3.78:1 at 3° and 2.31:1 at 7° retinal eccentricity in humans [66], and 5.43:1 at 6.4° and 4.08:1 at 7.5° in macaque [67]. In comparison, histologic human data from Gao & Hollyfield [68] reported cone-to-RPE ratios of 24.09:1 at the fovea and 0.89:1 at the equator. Panda-Jonas et al. [69] found an average ratio of 1.1 across the temporal meridian, spanning approximately 6.7° to 16.7°. Our measurement at 19.5° is consistent with these trends in the peripheral retina and highlights the extended imaging capability enabled by the AO-OCT system with the new fixation-stimulus channel.

Additionally at 19.5° retinal eccentricity, structural AO-OCT imaging of the vitreous-ILM border enabled visualization of individual macrophage-like cells [15,70,71], which are known to play important roles in immune responses [72]. Within the retinal ganglion cell layer, a monolayer of somas (comprising retinal ganglion cells and displaced amacrine cells) was clearly visible, along with retinal vasculature and nerve fiber bundles. In the inner nuclear layer, we observed a single bright, and large soma located at the INL border, consistent with previous reports of displaced retinal ganglion cells [73] (Fig. 6).

Functional AO-OCT imaging of cones at this same location (Fig. 7)—specifically using the cone ORG—was initially challenging due to the short cone outer segment length of ∼12 µm (Fig. 7(B), top). However, by adjusting the complex averaging window to accommodate the shorter outer segments observed at 19.5°, and by implementing a double-flash ORG protocol (which reduced the number of AO-OCT videos required), we demonstrated that S, M, and L cone spectral types could be classified both more rapidly and farther into the peripheral retina.

In this study we validated our fixation-stimulus channel exclusively on healthy subjects but have since used it to image subjects with color vision anomalies, retinal diseases, including glaucoma, retinitis pigmentosa, Stargardt disease, and models of age-related macular degeneration such as pentosan polysulfate sodium toxicity. In both healthy and diseased eyes, the fixation-stimulus channel has proven to be a valuable addition to the Indiana AO-OCT system, expanding our ability for examining retinal structure and function across a larger swathe of the retina.

5. Conclusion

We have developed and implemented an integrated fixation-stimulus channel for adaptive optics ophthalmoscopy that addresses several design challenges, including limited working distance, restricted steering field of view, and limited dioptric correction range. By leveraging primarily stock components—aside from a custom 3D-printed conic optical mount and a few generic machined parts—we achieved a cost-effective design that offers high performance in both fixation control and visible-light stimulation. This system supports precise and repeatable targeting of retinal regions, as well as delivery of temporally and spectrally controlled stimuli suitable for functional imaging. While our design was tailored for the Indiana AO-OCT system, the underlying principles, trade-offs, and component choices are broadly applicable to other AO ophthalmoscopic platforms. As such, this work contributes a practical and generalizable approach to enhancing the versatility and usability of AO imaging systems for studying retinal structure and function at the cellular level.

Supplemental information

Supplement 1. Describes our inner and outer retinal imaging and optoretinography protocols, with 3D renderings and part drawings of the fixation-stimulus system and a comprehensive parts list.

boe-16-11-4796-s001.pdf^{(3.3MB, pdf)}

Acknowledgment

We thank Physical Sciences Inc. for the original conic mount design and thank Dan Ferguson for cutting the dichroic mirror. We also thank Alessandra Carmichael-Martins, Thomas Gast, and Stephen Burns for assistance and use of their 3D printer for generating the first prototypes of our 3D conic mount. Also, a big thanks to Timothy Clark for machining the aluminum parts and adaptors.

Funding

National Eye Institute 10.13039/100000053 ( R01-EY018339, R01-EY029808, R01-EY035675).

Disclosures

D.T. Miller is a co-inventor on a U.S. patent that relates to AO-OCT. All other co-authors declare no conflicts of interest.

Data availability

Data underlying the results presented in this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.

Supplemental document

See Supplement 1^{(3.3MB, pdf)} for supporting content.

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Associated Data

This section collects any data citations, data availability statements, or supplementary materials included in this article.

Supplementary Materials

Supplement 1. Describes our inner and outer retinal imaging and optoretinography protocols, with 3D renderings and part drawings of the fixation-stimulus system and a comprehensive parts list.

boe-16-11-4796-s001.pdf^{(3.3MB, pdf)}

Data Availability Statement

Data underlying the results presented in this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.

[r1] 1.Liang J., Williams D. R., Miller D. T., “Supernormal vision and high-resolution retinal imaging through adaptive optics,” J. Opt. Soc. Am. A 14(11), 2884 (1997). 10.1364/JOSAA.14.002884 [DOI] [PubMed] [Google Scholar]

[r2] 2.Dubra A., Sulai Y., “Reflective afocal broadband adaptive optics scanning ophthalmoscope,” Biomed. Opt. Express 2(6), 1757 (2011). 10.1364/BOE.2.001757 [DOI] [PMC free article] [PubMed] [Google Scholar]

[r3] 3.Dubra A., Sulai Y., Norris J. L., et al. , “Noninvasive imaging of the human rod photoreceptor mosaic using a confocal adaptive optics scanning ophthalmoscope,” Biomed. Opt. Express 2(7), 1864 (2011). 10.1364/BOE.2.001864 [DOI] [PMC free article] [PubMed] [Google Scholar]

[r4] 4.Zou W., Qi X., Burns S. A., “Woofer-tweeter adaptive optics scanning laser ophthalmoscopic imaging based on Lagrange-multiplier damped least-squares algorithm,” Biomed. Opt. Express 2(7), 1986–2004 (2011). 10.1364/BOE.2.001986 [DOI] [PMC free article] [PubMed] [Google Scholar]

[r5] 5.Zou W., Burns S. A., “Testing of Lagrange multiplier damped least-squares control algorithm for woofer-tweeter adaptive optics,” Appl. Opt. 51(9), 1198–1208 (2012). 10.1364/AO.51.001198 [DOI] [PMC free article] [PubMed] [Google Scholar]

[r6] 6.Brunner E., Shatokhina J., Shirazi M. F., et al. , “Retinal adaptive optics imaging with a pyramid wavefront sensor,” Biomed. Opt. Express 12(10), 5969–5990 (2021). 10.1364/BOE.438915 [DOI] [PMC free article] [PubMed] [Google Scholar]

[r7] 7.Hampson K. M., Turcotte R., Miller D. T., et al. , “Adaptive optics for high-resolution imaging,” Nat. Rev. Methods Primers 1(1), 68 (2021). 10.1038/s43586-021-00066-7 [DOI] [PMC free article] [PubMed] [Google Scholar]

[r8] 8.Williams D. R., Burns S. A., Miller D. T., et al. , “Evolution of adaptive optics retinal imaging [Invited],” Biomed. Opt. Express 14(3), 1307 (2023). 10.1364/BOE.485371 [DOI] [PMC free article] [PubMed] [Google Scholar]

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PERMALINK

Integrated fixation and stimulus channel for adaptive optics ophthalmoscopy

Marcel T Bernucci

Yan Liu

Qiuzhi Ji

James A Crowell

Kristen Bowles-Johnson

Matthew J Keller

Donald T Miller

Abstract

1. Introduction

2. Methods

Fig. 1.

2.1. Fixation subsystem design

2.2. Badal optometer design considerations

Fig. 2.

2.3. Stimulus subsystem design

2.4. Considerations in selecting the multi-mode fiber and lenses in the stimulus subsystem

2.5. System validation, AO-OCT imaging, and optoretinography stimulation protocols

Fig. 3.

3. Results

Fig. 4.

Table 1. Summary of fixation-stimulus channel parameters.

Fig. 5.

Fig. 6.

Fig. 7.

4. Discussion

4.1. Fixation subsystem performance and key design trade-offs

Fig. 8.

4.2. Fixation subsystem comparison to literature

4.3. Stimulus subsystem performance and key tradeoffs

4.4. Structural and functional imaging of the peripheral human retina

5. Conclusion

Supplemental information

Acknowledgment

Funding

Disclosures

Data availability

Supplemental document

References

Associated Data

Supplementary Materials

Data Availability Statement

ACTIONS

PERMALINK

RESOURCES

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