2.1.

Full-Field OCT Setup

OCT is an optical technique using the principle of white light interferometry.¹¹ The setup is based on a Michelson interferometer utilizing a broadband light source where one mirror is replaced by the studied sample.¹⁷ Interference patterns are visible only if the optical path difference between both arms is smaller than the coherence length of the source. Low coherence length results in interference patterns with a restricted axial extension and consequently a precise location of the sample interfaces. Moreover, the use of a microscope objective performs an increase of the lateral resolution as in optical microscopy. In our case, two identical microscope objectives are included in each arm to obtain interference patterns when the beams are recombined, leading to a Linnik interferometer microscope.¹⁸ Water immersion objectives are chosen to reduce the light reflection from the surface of the sample and to avoid dispersion mismatch between both arms,¹⁹ especially in the case of an immerged sample.

Our experimental implementation is a full-field time domain OCT device (see Fig. 1 ). The light source is a halogen lamp coupled with a Köhler illumination for a homogeneous illumination. The incident light is split into two arms by a beamsplitter cube with a working range between 700 and $1100\mathrm{nm}$ . In both arms, mirrors oriented at a $45\text{-}\mathrm{deg}$ angle reflect light vertically to use immersion objectives with horizontal samples. Two identical water-immersion microscope objectives [ $\times 10$ , 0.3 numerical aperture (NA), Olympus] are used to provide good transverse resolution. A glass slab is placed in the reference arm and serves as a fixed reference mirror. The sample is mounted on a high-precision linear stage (M-714, Physik Instrument PI), which provides a maximum displacement of $7\mathrm{mm}$ . This shift is obviously much longer than the penetration depth because of the absorption and the scattering properties of the sample. After recombination, both beams are focused with a $30\text{-}\mathrm{mm}$ achromatic lens on the detector. The frames are recorded by a CCD camera, sensitive from $400\text{to}1000\mathrm{nm}$ , composed of a $659\text{-}\times 494\text{-pixel}$ matrix and that can reach $200\text{frames}∕\mathrm{s}$ (GC 660, Prosilica). To increase the sensitivity of the device, $2\text{-}\times 2\text{-pixel}$ binning is performed, reducing the number of effective pixels to $329\times 247$ .

Fig. 1

Schematic representation of the full-field OCT setup with an immerged cornea as the probed sample.

2.2.

Image Acquisition and System Characteristics

2.2.2.

Spectral sensitivity

The spectral transfer function of the setup was optimized for the study of biological samples, that is, in the near-IR range. Experimentally, this function mainly depends on both the camera spectral sensitivity and the emission spectrum of the light source. The halogen lamp is a thermal light source and its spectrum is centered in the IR spectral range and reaches the lowest wavelengths of the visible range when the intensity is high enough. The effective spectral sensitivity is the multiplication of both spectra and is limited by the halogen lamp at the lowest wavelengths and by the camera at the highest ones. The experimental spectral transfer function of our optical setup is calculated from the Fourier transform of the interferograms recorded on a glass sample, assuming that the Fresnel reflection coefficient of this material is independent of the wavelength (see Fig. 2 ). In our case, the spectral transfer function is centered at $810\mathrm{nm}$ and exhibits a $230\text{-}\mathrm{nm}$ full width at half maximum (FWHM). All the optical elements (beamsplitter, mirrors, microscope objectives, achromatic lens) used in the setup were chosen for their optimal optical properties in this spectral range.

Fig. 2

Spectral sensitivity of the FF-OCT setup

2.2.3.

Axial and transverse resolutions

The axial resolution $\Delta z$ (i.e., in the $z$ direction in the Fig. 1) depends on the source coherence length $l_c$ . Interference fringes are visible only if the optical path difference between both arms is smaller than the temporal coherence length of the source. The axial resolution is generally defined as the FWHM of the interferogram. If the source is assumed to have a Gaussian spectrum, the axial resolution, which depends on the mean wavelength ${\lambda }_0$ , on the FWHM $\Delta \lambda$ of the effective spectrum and on the refractive index of the medium $n$ , can be written as

$$\Delta z=\frac{l_c}{2}=\frac{2\ln 2}{n\pi }\frac{{\lambda }_0^2}{\Delta \lambda }.$$

In our case, the experimental axial resolution in water was measured to be $1\mu \mathrm{m}$ for a $0.95\text{-}\mu \mathrm{m}$ theoretical one (see Fig. 3 ). Note that this axial resolution is degraded at depths of several hundred micrometers due to dispersion mismatch in the two interferometer arms relative to the studied tissue in the sample arm.¹⁹

Fig. 3

Interferogram obtained from a glass slab (dotted line); each point corresponds to the recorded intensity as a function of the axial shift of the sample and the Gaussian shape modeled envelops (dashed line) with a FWHM equal to $1\mu \mathrm{m}$ which gives the experimental axial resolution.

The transverse resolution (i.e., in the $x$ and $y$ directions in Fig. 1) is limited by the diffraction. The Rayleigh criterion was used to determine the minimum resolvable detail. The theoretical resolution $r$ is then given by the numerical aperture of the objectives NA and the center wavelength of the effective spectrum ${\lambda }_0$ :

$$r=\frac{{\lambda }_0}{2\mathrm{NA}}.$$

In our case, the theoretical transverse resolution is $1.4\mu \mathrm{m}$ and it was experimentally measured to be $1.5\mu \mathrm{m}$ . The resolution is degraded when imaging is performed through biological tissues with complex features that induce a distortion of the optical wavefront and multiple scattering effects. The field of view of the camera is a $250\text{-}\times 190\text{-}\mu {\mathrm{m}}^2$ rectangle. To conclude, our experimental device enables 3-D imaging with a micrometer scale in the three directions.

2.3.

Human Cornea and Laser Surgery

Human corneoscleral rims were obtained from the eye bank (French Blood Center, Marseille, France). This study is based on a corpus of 10 human donor corneas to ensure representative results. All specimens were tissues rejected for transplantation based on standard operating protocols specified by the European Eye Bank Association, mainly due to their low endothelial cell density (below $2000\text{cells}∕{\mathrm{mm}}^2$ ), but these corneal grafts remain representative for the study of the cornea structure, especially for the stroma structure. The eyes were obtained within $36\mathrm{h}$ after death and stored in Optisol-GS. The storage in Optisol-GS does not exceed $10\text{days}$ . Corneas were delivered in CorneaMax (Eurobio, Les Ulys, France) medium. Some corneas were directly imaged, intact, immerged in balanced salt solution (BSS). The other corneas were used to study laser surgery effects in the corneal tissue. After completion of the cut, the corneal button was replaced in CorneaMax and OCT imaging was performed within $48\mathrm{h}$ .

Corneal incisions were performed using a Femtec laser (Tecnolas PerfectVision^©, Heidelberg, Germany). This apparatus consists of a pulsed solid-body (Nd:glass) laser with a repetition rate of $40\mathrm{kHz}$ that emits light with a wavelength of $1055\mathrm{nm}$ and a pulse duration of $500\text{to}800\mathrm{fs}$ . Laser beam energy of $3.2\text{to}3.4\mu \mathrm{J}$ and a spot separation of $3\mu \mathrm{m}$ were chosen. To cut the corneal specimen, corneoscleral discs were inserted in an artificial anterior chamber (Moria, Antony, France). Sodium hyaluronate 1.4% (Healon, AMO, Ettlingen, Germany) was instilled in the artificial anterior chamber as a viscoelastic up to a pressure of $20\mathrm{mm}$ Hg. Then, the corneas were mounted on the Femtec laser through an aplanation lens and centered using the reflex image of the diode lights of the laser. The aplanation lens provided by Tecnolas Perfect Vision is curved with a $10.5\text{-}\mathrm{mm}$ radius of curvature. The cutting process was performed from posterior to anterior (i.e., a circular movement of the laser beam was used to cut through endothelium, stroma forward to the corneal epithelium). Patterns of the corneal incisions were as follows. First, the laser was set to obtain a deep lamellar corneal cut (parallel to the corneal surface) set at $160\mu \mathrm{m}$ deep, then corneal vertical cuts were performed, i.e., the angle for the cut edge was set at $90\mathrm{deg}$ to achieve a cut line running perpendicular to the surface tangent. The cornea was then disconnected from the laser and the corneal button was replaced in CorneaMax for further examination. After removal of the sample from the initial solution and during the preparation of the tissue, the hydration of the corneal surface was maintained by instillation of BSS. Nevertheless, BSS on the surface is not used during laser experiments to avoid any excess of liquid that could interfere with the measurements. The sample preparation did not exceed $10\min$ . Afterward, it is crucial to minimize the experiment duration, to limit the dehydration of the corneal samples.

3.1.

Human Graft Corneas

Figure 4 shows a 3-D tomographic image of the entire thickness of the cornea with $1\text{-}\mu \mathrm{m}$ scanning steps. One of the axial cross sections of the previous volume of data is presented in Fig. 4. We can see that the epithelium (EP), Bowman’s layer (BM), stroma (ST), Descemet’s membrane (DM), and endothelium (E) from the anterior surface to the posterior surface of the cornea are clearly differentiated. These different layers can be easily identified as on routine histological corneal tissue sections [see Fig. 4]. The recorded signal level is quite high also from the deeper areas, and the posterior interface can be precisely located even if the light crosses $800\mu \mathrm{m}$ of corneal tissue. Furthermore, imaging quality is good despite this relatively great depth. This is mainly due to the weak dispersion mismatch in the two interferometer arms and to the transparency property of the cornea.

Fig. 4

(a) Three-dimensional tomographic image of the entire thickness of the cornea ( $250\times 190\times 900\mu {\mathrm{m}}^3$ stack, with $1\text{-}\mu \mathrm{m}$ axial steps) and (b) for a better view of the cornea structure a cross section of the 3-D OCT image is shown (along the dashed line). The different structures that comprised the cornea are indicated: epithelium (EP), Bowman’s membrane (BM), stroma (ST), Descemet’s membrane (DM), and endothelium (E), and (c) can be compared to a histological section of a healthy cornea.

Figure 4 shows a stacked organization with bright interfaces parallel to the surface of the cornea. This light reflection is due to the keratocytes in the corneal stroma. These cells are parallel to each other and the collagen lamellae are located between them. The stacked organization of the collagen lamellae can be identified on the axial tomographic image, because they appear between the keratocytes with a lower signal. In the posterior part, most of the collagen lamellae tended to be arranged parallel to the corneal surface. The organization revealed by the OCT appears remarkably similar to that of the histological section.

From the same 3-D tomographic data, cross sections parallel to the surface of the cornea can be realized (see Video 1 ). At around $10\mu \mathrm{m}$ deep, inside the epithelium, the epithelial cells and the nuclei boundaries are clearly visible on the transverse slices. Highly scattering structural details revealing the uppermost stromal keratocytes are visible at around $35\mu \mathrm{m}$ deep, just above the basement membrane. In this example, the low thickness of the epithelium is related to the manipulation for tissue preparation at the Eye Bank and before OCT imaging. The keratocytes are imaged along the entire thickness of the cornea and their number decreases from the anterior to the posterior stroma. Finally, the endothelial layer appears at around $800\mu \mathrm{m}$ deep. At this level, we can observe folds in Descemet’s membrane, observed as an arched structure. Hexagonal structures imaged at the level of endothelium correspond to endothelial cells.

Video 1

Two-dimensonal tomographic images along the transverse planes from the previous stack shown in Fig. 4 ( $250\times 190\times 900\mu {\mathrm{m}}^3$ stack, with $1\text{-}\mu \mathrm{m}$ axial steps); the scale bar is $50\mu \mathrm{m}$ and the depth is indicated in the upper left corner. The different structures that comprised the cornea are visible: epithelium (EP), Bowman’s membrane (BM), stroma (ST), Descemet’s membrane (DM), and endothelium (E). At 35 and $250\mu \mathrm{m}$ deep in the corneal stroma, the keratocytes are identified with their specific shape (QuickTime, $4.4\mathrm{MB}$ )..

At approximately $300\mu \mathrm{m}$ below Bowman’s layer (tomographic image in the lower left corner of Video 1), the keratocyte network is revealed as flat cells connected by cytoplasmic processes, linking adjacent cells together. Keratocytes give a high tomographic signal, even in a deep region of the stroma, and they can easily be recognized along the entire thickness of the cornea. Axial tomographic views [Fig. 4] are projections of the transverse views. The characteristic stacked organization is due to the tomographic signal of the keratocytes projected on a perpendicular plane.

In Fig. 5 , only the first $125\mu \mathrm{m}$ of the corneal structure are shown with an axial acquisition step reduced to $0.5\mu \mathrm{m}$ for a better imaging quality in the axial direction. The dynamic views of these tomographic data are presented on Video 2 . Intraepithelial layers [EP layer in Fig. 5 transverse view is presented in Fig. 5], Bowman’s layer [BM in Fig. 5], and anterior stroma [ST in Fig. 5] are clearly visualized from the anterior to the posterior part of the cornea. The resolution of the system is high enough to distinguish even intraepithelial layers and some epithelial wing cells appear at the top of the 3-D image. The transverse sections of the epithelium [Fig. 5] show the nucleis of the epithelial cells. The high-resolution images show a thickened epithelium with hyporeflective corneal epithelial bullae [black arrows in Fig. 5], corresponding to intraepithelial edema vacuoles. The position of the Bowman’s membrane can be clearly located [BM in Fig 5]. This membrane appears as a dark band with a thickness of $20\mu \mathrm{m}$ . The thin and curved layer oriented parallel to the corneal surface just above Bowman’s layer [see white arrow in Fig. 5] can be identified as the basement membrane. In Fig. 5, we could observe the anterior stroma. In this layer, the specific shape of the keratocytes is seen as a bright pattern surrounded by scattering collagen fibrils.

Fig. 5

(a) Three-dimensional tomographic image of the surface of the cornea ( $250\times 190\times 150\mu {\mathrm{m}}^3$ stack, with $0.5\text{-}\mu \mathrm{m}$ axial steps) with the epithelium (EP), the Bowman’s membrane (BM), the basement membrane (white arrow) and the anterior stroma (ST), and (b) and (c) 2-D tomographic images along the transverse direction respectively in the epithelium with the hyporeflective corneal epithelial bullae (black arrows) and in the stroma.

Video 2

Two-dimensional tomographic dynamic views of the upper part of the cornea ( $250\times 190\times 150\mu {\mathrm{m}}^3$ stack, with $0.5\text{-}\mu \mathrm{m}$ axial steps) along the different planes, successively transverse and axial. The epithelium (EP), the Bowman’s membrane (BM), and the anterior stroma (ST) are visible on these different views (QuickTime, $3.6\mathrm{MB}$ ). .

3.2.

Edematous Corneas

OCT imaging was also performed on edematous corneas (see Fig. 6 ) and enabled the observation of the water accumulation regions that are formed between stromal lamellae in strongly edematous corneas. The presence of water in the stroma appears as dark regions on the images because there is no light scattering in these regions and, hence, no tomographic signal is recorded. In these pathologic conditions, the degree of organization of the collagen fibers is decreased and the amount of interstitial fluid is increased. Large fluid vacuoles are present between lamellae, the so-called “lakes,”³⁹ resulting in distortion of fibrillar collagen distribution. On nonedematous cornea grafts, keratocytes are the main cells observed in the stroma with OCT imaging, whereas on an edematous cornea stromal lamellae become clearly visible [see Figs. 6 and 6], mainly due to the refraction index mismatch between the lakes, composed of water, and the collagen lamellae.

Fig. 6

Tomographic images ( $250\times 190\times 450\mu {\mathrm{m}}^3$ stack, with $0.5\text{-}\mu \mathrm{m}$ axial steps) along (a) the axial and (b) and (c) the transverse directions at 65 and $90\mu \mathrm{m}$ deep of an edematous cornea.

In parallel to these physiological changes, it is well known that the swelling of the cornea is correlated with a clinical reduction of the transparency of the tissue and a modification of its optical properties.^{40, 41} The observation of the growth of such edema is then very interesting for a good understanding of these mechanisms.

3.3.

Laser Surgery

The effect of femtosecond laser incisions in clear corneal tissue, at threshold exposure and above, can also be observed by ultrahigh-resolution OCT. Lamellar incisions were performed in the anterior stroma of human cornea at $160\mu \mathrm{m}$ deep (see Video 3 ). In this geometry, laser cuts are parallel to the corneal surface and show a slightly irregular cut pattern. The cutting edge inside the stroma is easily visible. The perturbation of the collagen structure extends to only a few micrometers of depth in the tissue and a hyperdensification of the collagen is visible at the borders of the incision, corresponding to disordering and delamination effects. Normal adjacent collagen fibers without any evidence of damage could also be observed. The laser cuts induced close to the threshold exhibited only localized corneal tissue modifications. Moreover, some cavitation bubbles remain $36\mathrm{h}$ after the laser incisions that are much smaller than the ones imaged just after the surgery, a few tens versus a few hundreds of micrometers. The size variations can be explained by the cavitation bubbles resorption within the first hours after laser surgery.

Video 3

Two-dimensional tomographic images along the axial direction; arrows indicate the microcavities due to the laser incisions at $160\mu \mathrm{m}$ deep in the stroma $36\mathrm{h}$ after laser surgery; scale bar is $50\mu \mathrm{m}$ ( $250\times 190\times 350\mu {\mathrm{m}}^3$ stack, with $0.5\text{-}\mu \mathrm{m}$ axial steps) (QuickTime, $1.1\mathrm{MB}$ ). .

Perforating femtosecond laser incisions (see Video 4 ) were investigated and we observe that the laser produces a regular incision of constant good quality across the depth of the cornea. A perturbation of the collagen structure as in lamellar incisions can be observed and irregular residual cavities remain present in the corneal tissue with dimensions up to several tens of micrometers. Obviously, although the incident pulse energy is kept constant, the laser beam broadening and attenuation due to the opacity of the tissue result in stronger disruption effects in the anterior part of the cornea than in the posterior region. At the surface of the vertical side cuts, OCT imaging seems to show a more reflective interface (arrow in Video 4) than on the other parts of the cornea surface. This could be due to the laser incisions in the upper layer of the cornea. This hyperreflective interface and the modification of the stroma structure with the presence of cavitation bubbles could induce the lack of signal from the deeper structures in this region observed in the lower left corner of the images. For OCT imaging, the cornea is illuminated from the anterior part (according to the vertical direction on the presented images) and the previous modifications of the structure could explain the shadow because the propagation of the light beam is modified and, hence, the penetration depth is smaller in this region.

Video 4

Two-dimensional tomographic images along the axial direction where the perforating laser incisions have been realized; scale bar is $50\mu \mathrm{m}$ ( $250\times 190\times 400\mu {\mathrm{m}}^3$ stack, with $0.5\text{-}\mu \mathrm{m}$ axial steps). The laser incisions on the surface on the cornea (arrow) show a hyperreflective structure (QuickTime, $2.5\mathrm{MB}$ ). .