# Transmission interference microscopy of anterior human eye

Samer Alhaddad, Wajdene Ghouali, Christophe Baudouin, Albert Claude Boccara, Viacheslav Mazlin

► To cite this version:

Samer Alhaddad, Wajdene Ghouali, Christophe Baudouin, Albert Claude Boccara, Viacheslav Mazlin. Transmission interference microscopy of anterior human eye. 2025. hal-04868627

HAL Id: hal-04868627

https://hal.science/hal-04868627v1

Preprint submitted on 6 Jan 2025

HAL is a multi-disciplinary open access
archive for the deposit and dissemination of sci-
entific research documents, whether they are pub-
lished or not. The documents may come from
teaching and research institutions in France or
abroad, or from public or private research centers.

L'archive ouverte pluridisciplinaire HAL, est
destinée au dépôt et à la diffusion de documents
scientifiques de niveau recherche, publiés ou non,
émanant des établissements d'enseignement et de
recherche français ou étrangers, des laboratoires
publics ou privés.

<figure><img src="image_1.png" /></figure>

Distributed under a Creative Commons Attribution 4.0 International License
# Transmission interference microscopy of anterior human eye

Samer Alhaddad<fnref n="1" />,<fnref n="2" />, Wajdene Ghouali<fnref n="3" />, Christophe Baudouin<fnref n="3" />, Albert Claude Boccara<fnref n="1" /> and Viacheslav Mazlin<fnref n="1" />,*

## Abstract:

We demonstrate that the fundamental ideas of E. Abbe about the interference nature of transmission microscopy are relevant for imaging of in vivo human eye. The transmission geometry in the eye is achieved by projecting illumination onto the posterior eye (sclera) and using the back-reflected light as a secondary illumination source for anterior eye structures. In particular, we show that the tightly localized illumination spot at the sclera acts as a closed condenser aperture in conventional microscopy, significantly enhancing interference contrast. In vivo cells and nerves are revealed with increased contrast across all corneal layers and within an extended 2 mm field of view, 7× larger than the clinical state-of-the-art. This transmission microscope presents the only modality, where the crystalline lens epithelial cells, fibers and sutures can be readily counted. In patients, Fuch's endothelial dystrophy, a major ophthalmic disease affecting 300 million people, is highlighted under a new transmission contrast. Finally, the instrument is built using consumer cameras, which opens opportunities for its wide adoption in the pre-screening and follow-up of cataract/refractive surgeries and the diagnosis of corneal infections in developing countries, where anterior eye diseases are most prevalent.

## Introduction:

Anterior segment, composed of cornea and crystalline lens, ranks among the most commonly operated parts of the human body. Every year, 26 million people worldwide undergo cataract surgery of the crystalline lens<fnref n="1" />, and 5 million people receive refractive surgery of the cornea<fnref n="2" />. The demand for surgeries is even higher with 65 million people awaiting for cataract surgery and 124 million people being visually impaired or blind due to refractive errors<fnref n="3" />. The surgeries have very high success rates of above 90%; nevertheless, the remaining 10% of cases still represent millions of people. The post-surgical complications range from dry eye disease (DED)⁴, resulting in the life-long ocular pain and the need to use the artificial tears, to endothelial decompensation necessitating the corneal transplantation⁵. Importantly, many undesired outcomes can be prevented through pre-surgical diagnosis. For example, a high-resolution specular microscope (SM) is used routinely in clinics to count corneal endothelial cell density and quantitatively assess the risk of potential decompensation⁶. As a drawback, SM is narrowly specialized in imaging the endothelial layer and is incapable of visualizing other corneal structures such as nerves that play a key role in DED.

Over the years, other instruments have been developed that enabled the visualization of the entire cornea with high resolution. Notably, in vivo confocal microscopy (CM) has become part of hundreds of clinical trials and has identified new cellular biomarkers for DED, corneal inflammation, infections, dystrophies⁷,⁸. CM has also become the first instrument to establish the connection between corneal biomarkers and general health conditions, e.g. diabetes⁹ and neurological conditions including Parkinson's disease and multiple sclerosis¹⁰,¹¹. Unfortunately, CM has remained complicated to use in daily clinics because the instrument required direct physical contact with the patient's eye and offered a very limited field of view (FOV) below 0.5 mm. More recently, those limitations of CM have been resolved by a new generation of high-resolution optical coherence tomography (OCT) devices that combined cellular resolution with an extended FOV of about 1 mm and non-contact imaging at a safe distance from the eye¹²⁻²². However, the elevated cost and complexity of high-resolution OCT components such as ultra-fast scientific cameras, secondary OCT sub-systems or high-frequency 2D galvo-scanning systems remained a significant barrier to their clinical transfer.

All the above-mentioned technologies operate in reflection (back-scattering geometry); however, it is also possible to image the human eye in transmission. Since the 1970s, ophthalmologists have been familiar with the

<footnote n="1">¹Institut Langevin, ESPCI Paris, Université PSL, Sorbonne Université, Université Paris Cité, CNRS, 75005 Paris, France</footnote>

<footnote n="2">²SharpEye SAS, France</footnote>

<footnote n="3">³Quinze-Vingts National Ophthalmology Hospital, 28 rue de Charenton, 75012 Paris, France</footnote>
slit-lamp retroillumination exam, where the back-scattered light from the posterior eye is used to illuminate the anterior eye structures²³. This back-scattered light is also familiar to photographers as the 'red eye' effect. Although useful for global exploration of the eye, this generic transmission design produced low-resolution and weak contrast, insufficient for resolving cellular details. One clever approach to see the cells in transmission was recently proposed by T. Weber and J. Mertz²⁴, who induced phase-gradient contrast²⁵ in the cornea and crystalline lens using asymmetric large-area illumination (42° visual angle) back-scattered from the posterior eye.

In this work we show that the ideas of E. Abbe²⁶ and F. Zernike²⁷ about the interference nature of transmission imaging are also relevant to the microscopy of the eye. In particular, the interference contrast can be enhanced by using symmetric illumination focused on a small localized spot (6°) on the posterior eye. As demonstrated below, the localized spot on the retina/sclera acts analogously to an almost closed condenser aperture (aperture controlling the NA of sample illumination) in conventional microscopy and increases the coherent contrast at the expense of reduction in resolution/depth-of-field (DOF). We demonstrate a significant enhancement in the visibility of corneal and crystalline lens structures in the human eye in vivo. Both healthy subjects and those with pathologies are examined, and the imaging results are compared with established clinical modalities (CM, SM) as well as an emerging research time-domain full-field OCT (TD-FF-OCT) device. The effects of the pupil size and cataract scattering on transmission imaging are investigated.

## Results:

### Physical concept of transmission interference microscopy in the eye

The instrument is shown in Fig. 1d. The light from the near-infrared (NIR) 850 nm light-emitting diode (LED) is focused at the back-focal plane of the microscope objective. The objective produces a quasi-collimated beam that is directed towards the eye and focused by the optics of the eye (cornea and crystalline lens) onto the retina. The sclera, located immediately behind the retina, scatters the incoming illumination and a portion of the scattered light returns to illuminate the cornea and crystalline lens from the back. The light passing through those layers in transmission is collected by the camera. Finally, the camera produces an image of a single anterior eye layer that is located in the optical focus of the objective. To capture only the transmitted light and avoid the strong specular reflex from the corneal surface, the illumination and detection are cross-polarized. Here, we rely on multiple scattering in the sclera to randomize the pre-encoded linear polarization of the illumination. The optical components are listed in *Methods*, while the photo of the prototype and its performance with a USAF target are shown in *Supplementary Fig. 1*.

Curiously, the imaged structures in this transmission microscope appear either with dark or bright contrasts depending on their axial location relative to the optical focus (Figs. 1a-c). This effect has been known in bright field transmission microscopy and is attributed to the interference nature of light²⁶. The interference occurs between the light diffracted by the sample and the light transmitted through the sample (zero-order diffraction). Analytically, it can be described by calculating the diffraction integrals at the sample, represented as a diffraction grating²⁸⁻³¹, and at the objective aperture, or simply by considering the Gouy phase shift in the Gaussian beam model³².

The insightful visualization of the above phenomenon comes from a simulation of the wave equation in Figs. 1e-i. We modelled the transmission as a 4F system with two lenses that conjugate the sample to the camera plane. For simplicity, a single spherical scatterer in the sample is illuminated from below with a quasi-collimated wave of incoherent light. See the simulation details in *Methods*. In the absence of the scatterer, the quasi-collimated wave is simply relayed by the 4F optical system, producing a uniform illumination on the camera. In contrast, when the scatterer is placed near the focal plane of the lens, a portion of the quasi-collimated wave close to the scatter is diffracted (diffracted wave), while another portion of the wave located further from the scatterer propagates without diffraction (transmitted wave). Although the two waves propagate within the same optical path they exhibit different angular divergences – the diffracted wave fills a significantly larger portion of the NA of the first lens. This leads to a spatial separation of the waves in the back-focal plane of the lens, where the transmitted wave is focused, and the diffracted wave is collimated (see the zoomed Figs. 1e,h and *Supplementary Video 1*). The second lens recombines both waves on the camera, producing dark or bright interference contrast for the scatterer. The thicker the scatterer and the higher its refractive index relatively to the background, the
larger the phase delay between the waves, leading to higher contrast. The sign of the contrast, denoting constructive or destructive interference, can be alternated by moving the camera, the scatterer, or the lenses.
For a small scatterer, a shift about the DOF is sufficient to change the contrast from the brightest to the darkest
(Fig. 1i).

<figure><img src="image_2.png" /><figcaption>Fig. 1. Principle of transmission interference microscopy in the eye. a, Measured intensity around the focal plane. The optical phase exhibits the characteristic π-phase shift. b,c, Transmission images from artificial eye used to compute the curve in a. d, Optical scheme of eye imaging device. f,g,i, Wave propagation simulations of equivalent 4F transmission imaging in the cases: without the scatter (f), and with the scatterer (g). The images are produced by time-averaging. e,h, The zoomed regions are produced without time-averaging and illustrate the separation of the focused transmitted wave (blue arrow) and collimated scattered wave (red arrows) in the back focal plane. The video of coherent wave propagation can be found in *Supplementary Video 1*. i, The shift of the scatterer around the focal plane changes its visible brightness in the camera plane, in agreement with the experiment. j, The retroilluminated light from the back of the eye, as seen through the imaging system using small magnification optics instead of the microscope objective.</figcaption></figure>
## Enhancement of interference contrast in the eye

The interference contrast is valuable as it enables label-free staining of transparent ocular tissues that would be invisible otherwise. One limitation of this contrast is that it is inherently weak, owing to the small phase retardation between the transmitted and diffracted waves. Since at least the early 20th century, it has been known that the visibility of scatterers in transmission microscopy can be controlled using the condenser aperture of the source³³. By closing the aperture one can reduce the NA illuminating the sample and significantly improve the interference contrast. We show that the same principle holds true for transmission imaging in the eye. In particular, the size of the light spot projected onto the sclera at the back of the eye affects the NA of the secondary source illuminating the anterior eye structures from the back and, consequently, determines the contrast.

A transmission microscope with interchangeable lenses and an aperture in the illumination arm was used to control the magnification of the source image projected onto the sclera (see Fig. 2a and also *Supplementary Fig.* 2a for the full optical scheme). For proper comparison, we imaged the same corneal location of the artificial model eye in all illumination configurations. See the details of the optical configurations in *Methods*. We observed a drastic increase in interference contrast when the size of the projected source was reduced (Fig. 2b,d). This tendency remained consistent across the full range of source magnifications tested, from 0.3 mm to 15 mm. On the flip side, the increase in contrast was accompanied by a decrease in DOF, visible as elongation of the axial phase profiles. Importantly, these changes were observed by adapting the illumination only, while the detection path remained intact.

The key idea for understanding these effects is that the source at the sclera (posterior eye) is always located at a distance equal to the length of the eye (about 23 mm) further from the image plane (anterior eye). As a result, the NA of the secondary illumination visible to the scatterer in the anterior eye is determined by the size of the source as: $NA_i \approx n_{\text{cornea-lens}} \cdot \frac{\text{Source radius}}{\text{Eye length}}$. The $NA_i$ affects both the lateral resolution and DOF, which are commonly determined from the algebraic mean of the $NA_i$ of illumination and $NA_d$ of detection³⁴: $\Delta x = \frac{0.61\lambda}{NA_{\text{eff}}}$ and $DOF = \frac{n \cdot \lambda}{NA_{\text{eff}}^2}$, where $NA_{\text{eff}} = \frac{NA_i + NA_d}{2}$ and n is the refractive index in the anterior eye. Indeed, experimentally, the scatterers illuminated with a small source (0.3 mm) and small $NA_i$ ($NA_i = 0.01$) fade significantly slower with defocus than the same scatterers illuminated with a large source (15 mm) and large $NA_i$ ($NA_i = 0.4$) (Fig. 2b).

The improvement of contrast with a smaller source size is explained with the simulation in Fig. 2c. When the source is large, the scatterer is exposed to a broad range of illumination angles. Each illumination angle produces a dark-bright phase profile in the camera plane that is tilted relatively to the optical axis. After summing up all these profiles, the contrast is noticeably reduced. Conversely, with a point source, the scatterer is illuminated by light of higher coherence, originating from a single illumination angle. This results in a more defined phase relationship between the transmitted and diffracted waves, leading to higher contrast.

In the clinical research prototype the illumination was designed with the intention to enhance the interference contrast over the resolution. The contrast is higher for a smaller illumination spot on the sclera. However, the conservation of *étendue* for an LED source sets the lower bound on the minimal spot size that can be achieved without losing light. Using a microscope objective with 0.3 NA, we produced a 1:1 projection of the LED chip onto the sclera resulting in a small 1.7 mm illumination field. Importantly, one has to also consider the enlargement of the spot size due to the diffuse scattering of the sclera. Using the mean free path of the sclera³⁵,³⁶ it is possible to estimate the size of the enlarged spot, which in our case is 2.4 mm (see *Methods*). Using this size, we can calculate the lateral resolution/DOF of 2.8 µm/34 µm. Based on Fig. 2d, one can deduce that this illumination is expected to provide a 3x gain in interference contrast compared to the other known retroillumination microscopes with extended (12 mm = 40°) illuminations²⁴,³⁷. We also tested the high-resolution configuration with a microscope objective of 0.4 NA. The 3 mm illumination spot was diffused to about 3.4 mm, resulting in a 2 µm/19 µm lateral resolution/DOF, respectively.
<figure><img src="image_3.png" /><figcaption>Fig. 2. Control of contrast and resolution in transmission interference microscopy. a, Reduced optical diagram of the experiment. The full optical scheme that shows the lenses and iris, used to adapt the size of the projected spot to the back of the artificial eye, can be found in *Supplementary Fig. 2a*. b, Left column: Different projected source sizes, measured experimentally by replacing the posterior part of the artificial eye with the camera. Middle column: En face transmission microscopy images from about 0.5 mm deep inside the cornea (denoted as 0 µm) and from about 0.58 mm deeper layer (denoted as 80 µm). Yellow arrows illustrate faster defocus of the dark spot (smaller DOF), when the illumination spot projected at the back of the eye is smaller. Right column: Axial cross-sectional images produced by stacking the zoomed en face images of the scatter. The 2D profile of the scatter highlights the greater interference phase shift contrast with smaller source sizes. The 3D stacks for different source sizes are visualized in *Supplementary Video 2*. c, Simulation of reduction in contrast and increase in DOF following an increase in the source size at the back of the eye. d, Quantitative dependency of contrast on the diameter of the illumination spot measured from b by calculating the difference between the maximal and minimal values of the axial profile for each spot size.</figcaption></figure>

Here, it is important to note that the formula for the $NA_{\text{eff}}$ that we used to estimate resolutions is known for being an approximation$^{38}$, and is, in fact, dependent on the scattering within the sample. Indeed, according to Mie theory, the small isolated particles with the size comparable to the wavelength of light exhibit isotropic forward scattering that fills the detection $NA_d$ independently of illumination $NA_i$, so that $NA_{\text{eff}} = NA_d^{39,40}$. This means that, in transmission, the isolated smallest particles could be seen with a better resolution determined by $NA_d$, independently of $NA_i$. On the other hand, larger particles have predominantly forward scattering that keeps the angular divergence of the illuminating light $NA_{\text{eff}} \approx NA_i$. The latter means that the larger particles can hinder viewing across an extended DOF determined by $NA_i$. Finally, the periodicity in the organization of the scatterers also determines the diffraction angles and $NA_{\text{eff}}$, similar to the effect seen with a diffraction grating.
Effects of pupil size and cataract scattering on the interference contrast

Reduced pupil size and increased scattering of the crystalline lens (cataract) are common among a large
population of elderly subjects. To study their effects on interference contrast, we used an artificial model eye
having replaceable anterior part featuring either dilated/contracted pupil or transparent/scattering lens. The
specification of the artificial eye is provided in *Methods*, while results are summarized in *Supplementary Fig. 2b,c*.

In vivo imaging of human subjects

For deployment in clinics, we constructed a compact (30cm × 30cm × 40cm) tabletop prototype (Supplementary Fig. 1a). It included the optical setup shown in Fig. 1d, mounted on a manual joystick stage for XYZ alignment to the eye, as well as the electronic controllers and a mini-PC. We tested two budget astrophotography cameras, both with high quantum efficiency in NIR: ASI432 (ZWO, China) with global shutter for undistorted high-resolution imaging and ASI585MC (ZWO, China) with increased number of pixels, allowing for extended FOV. See the full list of equipment with specifications in *Methods*.

Four healthy subjects and two subjects with anterior eye abnormalities were imaged. The images directly contained interference signals and only minimal processing was required. For instance, the processing of images captured with the ASI585MC camera was limited to high-pass filtering to remove the uniform light background.

Corneal epithelial cells were visible across an extended 1 mm FOV (Fig. 3a). We estimated the signal-to-noise ratio of 35dB in Fig. 3a by comparing the signal in the cornea with the standard deviation of noise in the corner of the image, outside of the eye. Despite the supposedly limited axial resolution, we could distinctly identify the superficial, wing and basal types of cells located at different layers within the 50 µm thick epithelium (Fig. 3b,c). The average cell diameters of the corresponding layers measured 40 µm, 20 µm and 10 µm, respectively. By implementing an additional focus tuneable lens in the detection, it was possible to acquire a fly-though volume of the epithelium (see *Supplementary Video 3* and *Methods*).

Immediately below the epithelium, we observed the sub-basal nerves as well as the dendritic cells (Fig. 3d). Nerve density biomarker could be measured in a semi-automatic way by segmenting nerves using a custom U-net neural network (NN) trained on our transmission data. A segmented image with the full FOV is provided as *Supplementary Fig. 3* and the NN architecture is mentioned in *Methods*. The measured density of 22 mm/mm² was in agreement with the state-of-the-art CM results from the same subject (Fig. 3f), as well as with values reported in the literature⁴¹. Compared to CM, the new instrument offered additional advantages, such as contactless imaging and 10× larger viewing area.

Half-millimetre deep inside the cornea, we captured endothelial cell layer (Fig. 4a). The cells were visible with bright or dark interference contrast, depending on their axial location relative to the optical focus. Due to the natural corneal curvature, both contrasts could be observed within the same image. We measured a corneal endothelial cell density of 3200 cells/mm², which was within the healthy range according to the literature⁴².

The first clinical patient was a 68-year-old male diagnosed with Fuch's endothelial dystrophy (FED). This patient also had a clinically confirmed cataract and was enlisted for provisional cataract surgery. The characteristic guttae of FED were visible and appeared to be composed of three zones: external 30-100 µm dark halo, 10-20 µm black rings, and 4-8 µm central brighter spot inside them (Fig. 4d). Given that the intensity in transmission microscopy is directly linked to the phase, this suggests a non-uniform internal structure of the guttae in terms of thickness and/or refractive index. Moreover, the guttae exhibited dark-bright shifts across the focus further confirming that they are primarily 'phase' and not light absorbing objects. In comparison, reflection methods, such as SM and full-field OCT depict guttae in a rather limited way, showing them as dark spots (Figs. e,f). Moreover, new transmission design gives access to 2 mm x 1 mm FOV resulting in 16x larger viewing area than the state-of-the-art SM. That being said, the reflection methods are fundamentally better at highlighting the cell edges, which are important clinical biomarkers. Indeed, cell edges are typically tilted compared to the cell facets and can redirect the light outside of the detection NA<sub>d</sub>, thus making the edges seen as dark in reflection. In contrast, the transmission is not sensitive to the tilt angle but to optical thickness. Since and the optical thickness is minimal at the edges, the phase shift is small, and the transmission signal is weak. On the other hand, the nuclei and the main cell body are thicker and give strong signal in transmission.
<figure><img src="image_4.png" /><figcaption>Fig. 3. Anterior cornea of healthy human subjects in vivo. **a**, Corneal epithelium imaged with a large FOV captured in a single camera frame using 0.3 NA objective and ASI585 camera. **b,c**, Detailed look at the superficial, wing and basal epithelial layers provided by the 0.4 NA objective and ASI432 camera. **d**, Sub-basal nerve plexus with visible nerves, basal cells and dendritic cells (yellow arrows). **e**, Overlay of nerve image with NN segmentation. The full segmented FOV is provided as *Supplementary Fig. 3*. Even during short sub-second acquisition the nerves alternate their phase and their visible intensity between dark and bright (see *Supplementary Video 4*) due to the presence of the fast axial movements, known for the cornea⁴³. **f**, Comparison with confocal microscopy image obtained on the same subject. The quality and flatness of the image was dependent on the expertise of the clinical operator.</figcaption></figure>
<figure><img src="image_5.png" /><figcaption>Fig. 4. Endothelium corneal layer in health and disease. **a**, Endothelial cells of healthy subject were visible across the extended 2 mm FOV achieved using 0.3 NA objective and ASI585 camera. **b,c**, Transmission revealed the cell bodies with dark or bright interference contrast depending on their location relatively to the focal plane. **d**, Endothelium of Fuch's endothelial dystrophy patient with guttata. Guttata had non-uniform internal structure (red arrows, yellows arrows). **e**, Comparison view of endothelial guttata taken with the research TD-FF-OCT device in another subject. The interference fringes artifacts obstruct the view. **f**, Comparison with specular microscopy from another subject. All presented images above are single camera frames without averaging.</figcaption></figure>

Corneal stroma appears in transmission as relatively uniform (Fig. 5a) - a striking divergence from its appearance in reflection, where bright keratocyte cells are highlighted over the dark background (Fig. 5b). We hypothesize this difference to be caused by the unique organization of the stromal fibrils that is responsible for corneal transparency. Indeed, previous simulations have shown that mutual interference between light waves scattered by the fibrils preserves the wavefront⁴⁴. In our experiment, this implies that the diffracted wave has the same phase as the transmitted wave, and, therefore, no interference contrast can be observed. Although few structures can be seen in the healthy stroma, this may not be the case for pathological cases. For example, in an elderly 82-years-old male subject, the images revealed fibre bundles passing through the stroma corresponding to corneal microfolds (striae) related to aging process. In terms of semiology, such striae are interesting, as they could also be an important biomarker for the early detection of keratoconus disease⁴⁵, which is characterized by the mechanical deformation of the cornea.
Another strategy to visualise the stromal keratocytes is to use the tomographic configuration of the transmission microscope. Previously, it has been shown that by acquiring the two phase shifted images, one can partially reject the out-of-focus light and recover the tomographic view in thick transparent ex vivo samples³²,⁴⁶,⁴⁷. Here, we implemented a setup with two cameras having different levels of defocus in order to simultaneously capture the two optical phases. After subtracting the phase frames and enhancing the weak signal using NN⁴⁸, the resulting tomographic image successfully revealed the 15 µm stromal keratocyte nuclei (see Fig. 5g and *Supplementary Video 5*).

<figure><img src="image_6.png" /><figcaption>Fig. 5. Corneal stroma in transmission microscopy and tomography. a, Transmission microscopy view of the healthy corneal stroma with visible stromal nerve. b, Comparative reflection image from TD-FF-OCT is more contrasted and clearly shows nuclei of keratocytes. c, Stroma of elderly subject had visible stripes, hypothesized to be corneal microfolds (striae). The images a-c were produced by averaging 15 frames. d, The comparative reflection TD-FF-OCT image from the same subject. Striae are highlighted using red arrows. e, Principle of tomographic imaging using 2 cameras. f,g, Comparison between transmission interference microscopy (left) and tomography (right). Yellow lines indicate the groups of keratocyte nuclei. Blue arrow points at stromal nerve.</figcaption></figure>

Finally, the interference contrast enabled one of the most detailed looks into the microstructure of the crystalline lens in vivo (Fig. 6). The lens epithelium was composed of about 10-15 µm cells, which was in agreement with specular photography findings, reported using experimental device in 1980s⁴⁹. The cells were packed with a density of about 5000 cells/mm². The lens fibers were regularly packed and merged into the Y-suture, previously observed in vivo only with low-resolution OCT⁵⁰.
<figure><img src="image_7.png" /><figcaption>Fig. 6. Crystalline lens of the human subject. a, Epithelium layer. b, Fiber layer with visible sutures. The images are produced by averaging about 40 raw frames.</figcaption></figure>

Discussion

Fundamentally, this work demonstrates an interesting example of controlling the coherence of light that is being scattered from random biological media, such as the sclera, for the benefit of imaging. Additionally, it highlights that the concept of the interference contrast in transmission extends beyond thin specimens and is applicable to thick tissues within the eye.

One of the instrument's primary applications is in surgical support. Today, the endothelial cell density measured by SM serves as an established clinical biomarker for the pre-screening of candidates for cataract and refractive surgeries. The transmission interference microscope has potential to enhance the precision of this biomarker by bringing an extended FOV over which many more cells could be counted. Moreover, the extended FOV ensures that regions affected by the excessive guttae are not overlooked, enabling the identification and exclusion of patients with FED (300 million people affected⁵¹) from certain surgeries. Beyond the inclusion/exclusion criteria, the microscope could assist doctors in selecting the type of surgery best suited for the patient. For example, PRK surgery might be recommended over LASIK when the patient's cornea exhibits early symptoms of dystrophy, such as microfolds (striae) in the anterior cornea. Additionally, the high-resolution microscope can validate next generation of refractive surgery instruments that rely on the microstructuring of the cornea⁵².

Corneal infections are another major target for the clinical trials of the transmission microscope. This technology could eliminate the need for invasive corneal scrapping and contact CM⁵³, by offering a non-contact diagnostic approach that minimizes the risk of cross-contamination and is suitable for individuals with fragile corneas or children. Additionally, given the relatively low cost of its instrumentation, which relies on consumer cameras, and the robustness of its common-path interferometric design, this device presents opportunities for screening the population in the rural areas of developing countries, where infections are most prevalent.

A promising application for the transmission microscope lies in the assessment of corneal nerves, as there is currently no imaging modality that allows their efficient, reproducible and non-invasive analysis. This would be particularly relevant in pharmaceutical industry, where it can provide quantitative feedback on the efficacy of novel treatments. For example, the effect of nerve growth factor (NGF) medication against the dry eye disease, which affects 20% of the world's population⁵⁴, could be assessed by measuring nerve density before and after treatment⁵⁵. Additionally, nerve studies related to diabetes and brain-related diseases, which were previously conducted with CM⁹,¹¹, could be further facilitated with the transmission microscope, capable of capturing an extended FOV in a single short exposure and non-invasively. Moreover, the view of the crystalline lens offered by this technology could support the validation of breakthrough stem cell therapies developed for cataract surgeries⁵⁶.
Another compelling research direction is to examine the evolution of transmission images over time. In particular, advancing the work of CM on monitoring the activity of dendritic cells⁵⁷, could provide new insights into the vicious cycle of inflammation in dry eye disease⁵⁸.

The rapid evolution of NIR CMOS technology, propelled by the growing demand for security applications, is anticipated to bring further increase in sensitivity and number of pixels in the coming years. These advancements could expand the FOV of microscopy to a degree approaching that of macro instruments in clinics, such as OCT and slit-lamp.

## Methods

### Components of the prototype

The light from the NIR 850 nm LED chip (M850L3, Thorlabs, USA) was refocused onto the back-focal plane of the microscope objective with 1:1 magnification using a lens pair (LA1422-B, Thorlabs, USA). We used two interchangeable microscope objectives. For medium resolution we employed the air objective (LMPLN10XIR, Olympus, Japan) with 10× magnification, 0.3 NA, 18 mm focal length (FL), and 18 mm working distance (WD). For high-resolution imaging the air objective (plan Apo NIR, Mitutoyo, Japan) having 20× magnification, 0.4 NA, 10 mm FL, and 20 mm WD was utilized. The objectives converted the focused image of the LED at the back-focal plane into the quasi-collimated beam incident on the eye. The cornea and crystalline lens then focused the beam onto the retina and sclera at the back of the eye.

The size of the illumination spot on the retina/sclera was computed from the size of the LED chip and the focal lengths of the lenses, microscope objective, and the eye. Although the LED chip was 1 mm in size, the actual visible size of the LED was 1.8 mm due to the presence of the aspheric dome overlying the lens, which was confirmed using the ZEMAX ray data provided with the LED. Considering the common 17 mm value for the focal length of the eye, the spot size was $1.8 \text{ mm} \times 17 \text{ mm}/18 \text{ mm} = 1.7 \text{ mm}$ for an objective with 18 mm FL and $1.8 \text{ mm} \times 10 \text{ mm}/18 \text{ mm} = 3 \text{ mm}$ for an objective with 10 mm FL.

The light, scattered from the sclera, illuminates the anterior eye structures from the back and the transmitted light was focused by the lens doublet onto the camera. We tested two budget astrophotography cameras, both with high quantum efficiency (QE) > 30% in NIR: ASI432 (ZWO, China), equipped with an imx432 sensor (Sony, Japan), or ASI585 (ZWO, China), equipped with an imx585 sensor (Sony, Japan). Using a camera with QE was important because of the inherently weak light power originating from the sclera. Indeed, the sclera scatters light uniformly in all directions⁵⁹, meaning that only a small fraction of the scattered light can contribute to the back-illumination of the central cornea and eventually reach the detector. The ASI432 featured a global shutter and increased imaging speed of 120 frames/s (fps), which was beneficial for suppressing artefacts related to eye movement. The ASI585MC, on the other hand, had a rolling shutter sensor and a slower 45 fps rate, but it featured an increased number of pixels (3840 x 2160), allowing for an extended FOV. Prior to the usage protected window was removed from the camera to increase transmittance in NIR.

The focal length of the tube lens was adapted to the pre-selected objective and pixel size of the camera to achieve sampling that adhered to the Nyquist criterion. Specifically, for a 0.3 NA objective and ASI585, we used a 100 mm doublet (AC254-100-B, Thorlabs, USA) that oversampled the resolution.

To collect the transmitted light and avoid the strong specular reflection from the corneal surface the polarizing beam splitter (PBS252, Thorlabs, USA) was used. An additionally implemented linear polarizer (LPNIRE100-B) was necessary to produce a cleaner polarization state in the illumination. The optical setup was mounted on a clinical joystick stage capable of moving in X, Y, Z dimensions (purchased from Lumedica, USA). The optical setup could also include a focus-tuneable lens (EL-16-40-TC-NIR-5D-C, Optotune, Switzerland) implemented in the detection path between the beam splitter and the tube lens. By synchronizing the tuneable lens with the acquisition, the fly-through volume stack of the anterior cornea could be acquired in less than 1 s (Supplementary Video 3). For this, we also cropped the sensor area by half, increasing the frame rate of ASI432 camera to 178 fps.
The full device was controlled by an Arduino Uno board (Arduino.cc, Italy) and a mini-PC (BXNUC9i9QNX, Intel, USA). The Arduino managed the pulsating of the LED with 0.5 ms time precision. The mini-PC, with a custom-written Python code, collected 8-bit or 16-bit images from the cameras via the rapid USB 3.1 port, processed them, and displayed the result in real-time. To address the limitations of the budget cameras lacking a trigger feature, we implemented a software-based rolling buffer. Only the image with the highest intensity within the rolling buffer, captured in the moment when LED pulsation matched camera exposure, was displayed on the screen.

## Simulation of wave propagation

We deployed the scalar wave equation simulator from⁶⁰. Refractive indexes of the medium and lenses were 1 and 1.5, respectively. The scatterer had a size comparable to the wavelength and an elevated refractive index of 2.5 for more contrasted visualization. The 4F system contained identical lenses with the scatterer and the camera being close to the focal plane of the respective lenses. In order to reduce the coherent artifacts in the simulation, we generated the quasi-incoherent incident illumination by using a random assembly of particles (n=1.5, diameter about ¼ of the wavelength) that randomized the phase of the incoming wavefront in time. The Figs. 1f,g,i were time-averaged. By positioning the assembly closer or further from the scatterer, one could also control the divergence of the illumination angles to which the scatter was exposed.

## Experiment with variable illumination size

We used an artificial eye (Model Eye for OCT, Modell-Augen Manufaktur, Germany) having a modular structure. The anterior part of the eye was interchangeable and configured with either a 4 mm or 8 mm pupil and a transparent or light-scattering lens, the latter resembling a cataract condition. The sclera below the retina was made of white, solid scattering material that was translucent to light.

In the experiment with variable illumination size, the 0.3 NA objective and ASI585 camera were used. The magnification of the light spot at the sclera was adjusted by varying the focal lengths of the lenses or by using an aperture in the illumination arm. Sub-figures in Fig. 2b were acquired using the lenses pairs of 8 mm/50 mm, 16 mm/50 mm, 30 mm/50 mm and 50 mm/50 mm, respectively. In case of the smallest illumination size, we used 50 mm/50 mm lenses with an additional 200 µm aperture in front of the LED. The final size of the magnification spot was calculated using the formulas discussed above, with one modification: the artificial eye had a longer focal length of 25 mm due to a smaller corneal curvature. In the example of the 200 µm aperture the projected light spot was 0.2 × 25 mm/18 mm ≈ 0.3 mm, while in the case of 8 mm/50 mm it was 1.8 mm × 50 mm/8 mm × 25 mm/18 mm ≈ 15 mm. We confirmed the illumination sizes experimentally by replacing the posterior part of the eye with the second camera (ASI432).

The objective was focused on a single corneal plane of the artificial eye. In order to reconstruct the axial cross-section image and 3D volume, the artificial eye was moved axially using a motorized stage (X-DMQ12P-DE52, Zaber, Canada). To facilitate comparison, the axial intensity profiles were normalized between 0 and 1, based on the maximal and minimal intensity values measured in the most contrasted case of the 0.3 mm spot size. The relation between the contrast and illumination spot in Fig. 2d was obtained by calculating the difference between the maximal and minimal values of the axial profile for each spot size.

## Light diffusion in sclera

The secondary illumination source re-emitted from the sclera is larger than the spot size of the incident illumination due to light diffusion in the sclera. From the reduced scattering coefficient of 30 cm⁻¹ for the sclera³⁵,³⁶, the transport mean free path was computed as $l^* = 1/30 cm^{-1} = 0.33 mm$. Considering forward and backward propagation through the 1 mm thick sclera the increased spot size is given by:

$$D = 2 \cdot \sigma_{\text{total}} = 2 \cdot \sqrt{\sigma_0^2 + 2 \cdot z \cdot l^*}, \text{ where } \sigma_{\text{total}} \text{ is the total variance, } \sigma_0 \text{ is the initial spot size and } z \text{ is 1 mm sclera thickness.}$$

## Imaging of human subjects
The instrument was validated by imaging two male (32 y/o, 28 y/o) and two female (41 and 43 y/o) healthy subjects as well as one male patient (68 y/o) and one elderly subject (82 y/o) with corneal abnormalities. Prior to the research exam, all subjects were screened using the conventional slit-lamp in the hospital, which confirmed their corneal conditions. In particular, the patient displayed advanced Fuch's endothelial dystrophy and cataract. The elderly subject exhibited mild corneal haze typical for this age group. In addition to the exam with the new device, the elderly subject was imaged with two other cell-resolution modalities, including a clinical specular microscope (SP-3000P, Topcon, Japan) and a time-domain full-field OCT research instrument¹³. One healthy subject was also screened with confocal microscope (HRT3 RCM, Heidelberg engineering, Germany) and a full-field OCT. These high-resolution devices, which operate in reflection, served as a comparison reference for the new transmission images.

The subjects were asked to place their head on a clinical chinrest with forehead support. The examination was performed in a dark room to allow natural dilation of the pupil. No medications were introduced in the eye. In accordance with the French regulators, the approval for the study was obtained from Comité de Protection des Personnes (CPP) Sud-Ouest et Outre Mer III, study number 2024-A01040-47. Each subject has provided informed consent for participation in the study. Imaging was non-contact, with a comfortable 2 cm distance between the objective and the eye, and did not involve pupil dilation drugs. The instrument was emitting pulsed NIR illumination with radiant exposure adhering to the international light safety standards ISO 15004-2:2007 and ANSI Z80.36-2021. The light power was distributed across both time and space. More precisely, the 10 ms light pulse was followed by the 90 ms interval without light. At the retinal plane, the illumination was spread across the full-field area of 1.7 mm that was considerably larger than the micrometric focused spot, typical for scanning OCT devices. As a result, the total incident power of 100 mW produced a time-averaged weighted retinal irradiance of 100 mW × 0.01 s/(0.01 s + 0.09 s)/0.017 cm² × R ≈ 300 mW/cm², where R is the spectral weighting coefficient (R = 0.5 for 850 nm) according to ISO 15004-2:2007. The irradiance value was below the maximum permissible limit of 1.2/1.7 W/cm² = 700 mW/cm² (Table 4, ISO 15004-2). Similarly, the safety on the short time scale (Table 6, ISO 15004-2) was confirmed. For t = 20 s the weighted time-integrated radiant exposure was 100 mW/0.017 cm² × 20 s × 0.01 s/(0.01 s + 0.09 s) × R ≈ 6 J/cm², smaller than the permissible limit of 10/1.7 × (20 s)³/⁴ × (Nₚulses)⁻¹/⁴ J/cm² ≈ 14.5 J/cm². Here the Nₚulses is the number of pulses during 20 seconds. The exposure of the anterior eye was relatively weaker due to the collimation of the beam incident on the eye. Indeed, the measured irradiance of 5 mW/mm² at the corneal plane resulted in time-averaged irradiance of 50 mW/cm², below the 100 mW/cm² limit (Table 4, ISO 15004-2). On the short time scale of 20 s, the exposure of 1 J/cm² was smaller than the maximal permitted 3.8 J/cm² (Table 6, ISO 15004-2). We also verified the corneal and retinal safety for all the times below 20 s.

In an alternative mode, we illuminated the eye with a 1 s exposure followed by the 19 s interval without light. This mode was useful for acquiring more frames with minimal time delay, thus allowing to improve the SNR through image averaging. Repeating the calculations above with t = 1 s (the highest exposure condition) confirms that the time-integrated radiant exposures for retina of 3 J/cm² and cornea of 0.5 J/cm² were below their respective limits of 6 J/cm² and 1.8 J/cm².

## Image processing

The processing steps were dependent on the camera type. The ASI585 was a color camera, which required a gaussian blur with minimal kernel size (3 x 3 pixels) in order to uniformize the illumination mismatch between the different segments of the Bayer filter. Conversely, the ASI432 was a mono camera with a front-illuminated architecture, which introduced fixed pattern noise. This noise was removed by subtracting the average from the stack of moving images of the eye. Additionally, for both cameras we applied high pass filter with a large kernel (121 x 121 pixels) that suppressed the bright uniform background. In selected cases, for example in crystalline lens data, the images were averaged in order to increase SNR.

## Neural network architecture

We employed the U-net neural network architecture, as detailed in⁴⁸, for both segmentation and signal enhancement tasks. The segmentation network was trained using the in vivo transmission data from sub-basal nerve layer. Ground truth images were generated by segmenting the nerves in a semi-automated way using
NeuronJ plugins of Fiji⁶¹. The signal enhancement neural network, trained on full-field OCT data, was deployed from⁴⁸ and directly applied without retraining to enhance the weak tomographic transmission signals.

## Data availability

All images that support the findings of this study are presented in the manuscript.

## Code availability

The neural network code used in this article is available in Github repository: https://github.com/vmazlin/i2i.

## Acknowledgements

We would like to thank Daniel Royer for fruitful discussions.

## Author contributions

Idea and concept (V.M., S.A., A.C.B.), physical theory and simulations (V.M., S.A., A.C.B.), development of the device and experiments (V.M., S.A., A.C.B.), software for acquisition and image analysis (S.A., V.M), clinical guidance (W.G., C.B.), interpretation of biomarkers (W.G., C.B., V.M., S.A.), neural network processing (V.M., S.A.), safety analysis (V.M.), prototype 3D design and printing (S.A., V.M.). V.M. has written the manuscript, S.A. prepared the supplementary materials and all authors contributed with their feedback.

## Conflict of interest

V.M., S.A., A.C.B.: Patent; S.A.: Part-time salary from the company (SharpEye); V.M., A.C.B.: Ownership of the company shares (SharpEye)

## Funding

We would like to acknowledge the funding support of Agence Nationale de la Recherche grants: ANR-22-CE19-0018, ANR-22-CE45-0005, ANR-17-CONV-0005 (Q-life project TIOS).

## References

1. Chen, X., Xu, J., Chen, X. & Yao, K. Cataract: Advances in surgery and whether surgery remains the only treatment in future. *Adv. Ophthalmol. Pract. Res.* **1**, 100008 (2021).

2. Honavar, S. Refractive surgery - the 20/happy game changer. *Indian J. Ophthalmol.* **68**, 2639 (2020).

3. World Health Organization. *World Report on Vision*. (World Health Organization, Geneva, 2019).

4. Denoyer, A. et al. Dry Eye Disease after Refractive Surgery. *Ophthalmology* **122**, 669–676 (2015).

5. Feizi, S. Corneal endothelial cell dysfunction: etiologies and management. *Ther. Adv. Ophthalmol.* **10**, 2515841418815802 (2018).

6. Chaurasia, S. & Vanathi, M. Specular microscopy in clinical practice. *Indian J. Ophthalmol.* **69**, 517–524 (2021).

7. Guthoff, R. F., Baudouin, C. & Stave, J. *Atlas of Confocal Laser Scanning In-Vivo Microscopy in Ophthalmology*. (Springer Berlin Heidelberg, Berlin, Heidelberg, 2006). doi:10.1007/3-540-32707-X.

8. Villani, E. et al. *In Vivo Confocal Microscopy of the Ocular Surface: From Bench to Bedside*. *Curr. Eye Res.* **39**, 213–231 (2014).

9. Petropoulos, I. N. et al. Corneal Confocal Microscopy: A Biomarker for Diabetic Peripheral Neuropathy. *Clin. Ther.* **43**, 1457–1475 (2021).

10. Petropoulos, I. N. et al. Corneal confocal microscopy: ready for prime time. *Clin. Exp. Optom.* **103**, 265–277 (2020).

11. Petropoulos, I. N. et al. Corneal Confocal Microscopy to Image Small Nerve Fiber Degeneration: Ophthalmology Meets Neurology. *Front. Pain Res.* **2**, 725363 (2021).

12. Mazlin, V. et al. In vivo high resolution human corneal imaging using full-field optical coherence tomography. *Biomed. Opt. Express* **9**, 557–568 (2018).

13. Mazlin, V. et al. Real-time non-contact cellular imaging and angiography of human cornea and limbus with common-path full-field/SD OCT. *Nat. Commun.* **11**, 1868 (2020).

14. Aukorius, E. et al. In vivo imaging of the human cornea with high-speed and high-resolution Fourier-domain full-field optical coherence tomography. *Biomed. Opt. Express* **11**, 2849 (2020).

15. Tan, B. et al. 250 kHz, 15 µm resolution SD-OCT for in-vivo cellular imaging of the human cornea. *Biomed. Opt. Express* **9**, 6569 (2018).

16. Han, L. et al. Line-scanning SD-OCT for in-vivo, non-contact, volumetric, cellular resolution imaging of the human cornea and limbus. *Biomed. Opt. Express* **13**, 4007 (2022).

17. Pattan, H. F., Liu, X. & Tankam, P. Non-invasive in vivo imaging of human corneal microstructures with optical coherence microscopy. *Biomed. Opt. Express* **14**, 4888 (2023).

18. Neuhaus, K. et al. Real-time line-field optical coherence tomography for cellular resolution imaging of biological tissue. *Biomed. Opt. Express* **15**, 1059 (2024).

19. Chen, K., Song, W., Han, L. & Bizheva, K. Powell lens-based line-field spectral domain optical coherence tomography system for cellular resolution imaging of biological tissue. *Biomed. Opt. Express* **14**, 2003 (2023).
20. Chen, K., Abbasi, N., Wong, A. & Bizheva, K. In vivo, contactless, cellular resolution imaging of the human cornea with Powell lens based line field OCT. *Sci. Rep.* **14**, 22553 (2024).

21. Zhang, J. et al. Time-domain full-field optical coherence tomography (TD-FF-OCT) in ophthalmic imaging. *Ther. Adv. Chronic Dis.* **14**, 1197–1199 (2023).

22. Mazlin, V. et al. Curved-field optical coherence tomography: large-field imaging of human corneal cells and nerves. *Optica* **7**, 872 (2020).

23. Brown, N. Visibility of transparent objects in the eye by retroillumination. *Br. J. Ophthalmol.* **55**, 517–524 (1971).

24. Weber, T. D. & Mertz, J. In vivo corneal and lenticular microscopy with asymmetric fundus retroillumination. *Biomed. Opt. Express* **11**, 3263 (2020).

25. Ford, T. N., Chu, K. K. & Mertz, J. Phase-gradient microscopy in thick tissue with oblique back-illumination. *Nat. Methods* **9**, 1195–1197 (2012).

26. Abbe, E. A Contribution to the Theory of the Microscope and the Nature of Microscopic Vision. *Proceedings of the Bristol Naturalists' Society* 200–261 (1876).

27. Zernike, F. How I Discovered Phase Contrast. *Science* **121**, 345–349 (1955).

28. Born, M., Wolf, E. & Bhatia, A. B. *Principles of Optics: Electromagnetic Theory of Propagation, Interference, and Diffraction of Light*. (Cambridge University Press, Cambridge, 2019).

29. Françon, M. *Le Contraste de Phase En Optique et En Microscopie*. (Revue d'Optique, 1950).

30. Wade, R. H. A brief look at imaging and contrast transfer. *Ultramicroscopy* **46**, 145–156 (1992).

31. Hernández Candia, C. N. & Gutiérrez-Medina, B. Direct Imaging of Phase Objects Enables Conventional Deconvolution in Bright Field Light Microscopy. *PLoS ONE* **9**, e89106 (2014).

32. Mazlin, V., Thouvenin, O., Alhaddad, S., Boccara, M. & Boccara, C. Label free optical transmission tomography for biosystems: intracellular structures and dynamics. *Biomed. Opt. Express* **13**, 4190 (2022).

33. Dempster, W. T. Principles of Microscope Illumination and the Problem of Glare. *J. Opt. Soc. Am.* **34**, 695 (1944).

34. Murphy, D. B. & Davidson, M. W. *Fundamentals of Light Microscopy and Electronic Imaging*. (Wiley, 2012). doi:10.1002/9781118382905.

35. Bashkatov, A. N., Genina, E. A., Kochubey, V. I., Kamenskikh, T. G. & Tuchin, V. V. Optical clearing of human eye sclera. in (eds. Manns, F., Söderberg, P. G. & Ho, A.) 71631R (San Jose, CA, 2009). doi:10.1117/12.812394.

36. Rol, P., Niederer, P., Dürr, U., Henchoz, P. & Fankhauser, F. Experimental investigations on the light scattering properties of the human sclera. *Lasers Light Ophthalmol.* 201–221 (1990).

37. Meyer, P. A. R. Living histopathology - interrogation of ocular tissues by light: a celebration of the slit-lamp and a repertoire of clinical techniques. *Eye* (2024) doi:10.1038/s41433-024-03139-5.

38. Baker, J. R. Remarks on the Effect of the Aperture of the Condenser on Resolution by the Microscope. *J. Cell Sci.* **S3-93**, 375–378 (1952).

39. Taylor, R. W. & Sandoghdar, V. Interferometric Scattering Microscopy: Seeing Single Nanoparticles and Molecules via Rayleigh Scattering. *Nano Lett.* **19**, 4827–4835 (2019).

40. Boccara, M. et al. Full-field interferometry for counting and differentiating aquatic biotic nanoparticles: from laboratory to Tara Oceans. *Biomed. Opt. Express* **7**, 3736 (2016).

41. Parissi, M. et al. Standardized Baseline Human Corneal Subbasal Nerve Density for Clinical Investigations With Laser-Scanning in Vivo Confocal Microscopy. *Investig. Ophthalmology Vis. Sci.* **54**, 7091 (2013).

42. McCarey, B. E., Edelhauser, H. F. & Lynn, M. J. Review of Corneal Endothelial Specular Microscopy for FDA Clinical Trials of Refractive Procedures, Surgical Devices, and New Intraocular Drugs and Solutions. *Cornea* **27**, 1–16 (2008).

43. Mazlin, V. et al. Optical phase modulation by natural eye movements: application to time-domain FF-OCT image retrieval. *Biomed. Opt. Express* **13**, 902 (2022).

44. Meek, K. M. & Knupp, C. Corneal structure and transparency. *Prog. Retin. Eye Res.* **49**, 1–16 (2015).

45. Grieve, K. et al. Stromal striae: a new insight into corneal physiology and mechanics. *Sci. Rep.* **7**, 13584 (2017).

46. Alhaddad, S., Thouvenin, O., Boccara, M., Boccara, C. & Mazlin, V. Comparative analysis of full-field OCT and optical transmission tomography. *Biomed. Opt. Express* **14**, 4845 (2023).

47. Gutiérrez-Medina, B. Optical sectioning of unlabeled samples using bright-field microscopy. *Proc. Natl. Acad. Sci.* **119**, e2122937119 (2022).

48. Mazlin, V. Optical tomography in a single camera frame using fringe-encoded deep-learning full-field OCT. *Biomed. Opt. Express* **15**, 222 (2024).

49. Brown, N. A. P. & Bron, A. J. An estimate of the human lens epithelial cell size in vivo. *Exp. Eye Res.* **44**, 899–906 (1987).

50. Gupta, A. et al. In vivo SS-OCT imaging of crystalline lens sutures. *Biomed. Opt. Express* **11**, 5388 (2020).

51. Aiello, F., Gallo Afflitto, G., Ceccarelli, F., Cesareo, M. & Nucci, C. Global Prevalence of Fuchs Endothelial Corneal Dystrophy (FECD) in Adult Population: A Systematic Review and Meta-Analysis. *J. Ophthalmol.* **2022**, 1–7 (2022).

52. Linz, N., Freidank, S., Liang, X.-X. & Vogel, A. Laser Micro- and Nanostructuring for Refractive Eye Surgery. in *Ultrafast Laser Nanostructuring* (eds. Stoian, R. & Bonse, J.) vol. 239 1217–1245 (Springer International Publishing, Cham, 2023).

53. Labbé, A. et al. Contribution of In Vivo Confocal Microscopy to the Diagnosis and Management of Infectious Keratitis. *Ocul. Surf.* **7**, 41–52 (2009).

54. The Pathophysiology, Diagnosis, and Treatment of Dry Eye Disease (30.01.2015). https://www.aerzteblatt.de/int/archive/article/167472.

55. Sacchetti, M. et al. Effect of recombinant human nerve growth factor eye drops in patients with dry eye: a phase IIa, open label, multiple-dose study. *Br. J. Ophthalmol.* bjophthalmol-2018-312470 (2019) doi:10.1136/bjophthalmol-2018-312470.

56. Lin, H. et al. Lens regeneration using endogenous stem cells with gain of visual function. *Nature* **531**, 323–328 (2016).

57. Downie, L. E. et al. Redefining the human corneal immune compartment using dynamic intravital imaging. *Proc. Natl. Acad. Sci.* **120**, e2217795120 (2023).

58. Yamaguchi, T. Inflammatory Response in Dry Eye. *Investig. Ophthalmology Vis. Sci.* **59**, DES192 (2018).

59. Vogel, A., Dlugos, C., Nuffer, R. & Birngruber, R. Optical properties of human sclera, and their consequences for transscleral laser applications. *Lasers Surg. Med.* **11**, 331–340 (1991).

60. Morgensztern, E. starfree/wave-simulator. (2024).

61. Schindelin, J. et al. Fiji: an open-source platform for biological-image analysis. *Nat. Methods* **9**, 676–682 (2012).
Supplementary Figures:

<figure><img src="image_8.png" /><figcaption>a</figcaption></figure>

<figure><img src="image_9.png" /><figcaption>Supplementary Fig. 1. a, Tabletop clinical research prototype with a computer. b,c, Imaging of the USAF target in reflection and in transmission, where the scattering paper retroilluminates the target. Crossed polarization effectively suppresses the reflected light from the target while allowing the light scattered from the paper behind the target to reveal the latter in transmission.</figcaption></figure>

<figure><img src="image_10.png" /><figcaption>b</figcaption></figure>

<figure><img src="image_11.png" /><figcaption>c</figcaption></figure>

<figure><img src="image_12.png" /><figcaption>Supplementary Fig. 2. a, Full device from the experiments with the artificial eye. b, Effect of pupil size on the signal. The pupil size determines the width of the quasi-collimated illumination beam injected into the eye and, therefore, directly affects the amount of detected light on the camera. Although it is optically feasible to produce a narrow illumination beam that would not be partially blocked by the pupil, this implementation comes at the cost of reduced contrast, either due to increased source magnification or less efficient light collection from the LED. Therefore, pupil dilation should still be considered as a solution for improving images in cases where the signal from the eye is weak. c, Influence of crystalline lens scattering (cataract) on the interference contrast and on DOF/resolution. The scattering reduces interference contrast but also makes DOF shallower – a similar outcome to that expected by increasing the source size. Indeed, the scattering of the crystalline lens diffuses the illumination across the larger area of the sclera, thereby increasing the NA of the secondary illumination. Even more significant increase in NA occurs during the back-pass, when the light is getting scattered by the lens located in close proximity to the image plane. One solution to recover the contrast in cataract patients is to project a smaller source size compared to healthy subjects.</figcaption></figure>
<figure><img src="image_13.png" /><figcaption>Supplementary Fig. 3. Segmentation of sub-basal nerves using neural network.</figcaption></figure>