Light-induced polymerization and crosslinking are widely used methods for the fabrication and processing of polymeric materials, with applications ranging from manufacturing, photolithography, and tissue engineering. In most cases, the photochemical mechanism involves light absorption by a photosensitive compound, leading to the generation of free radicals that promote the polymerization of monomers or the crosslinking of polymers [1]. These methods are increasingly common for clinical applications since they enable the fabrication and manipulation of biomaterials in situ, including curing of dental composites [2] and photocrosslinking of polymer implants for regenerative medicine applications [3]. Some biological tissues, such as the cornea, are composed of intricate collagen fiber networks that can degrade due to diseases, but crosslinking can restore their natural biomechanical properties and functions [4,5]. Corneal ectatic disorders, for example, caused by the weakening of the corneal stroma, is effectively treated with photochemical crosslinking using riboflavin and ultraviolet A (UVA) light to strengthen corneal tissue [6].

While most applications rely on single-photon absorption (typically UV light), two-photon absorption-mediated processes using a near-infrared (NIR) femtosecond laser offers far superior spatial resolution that is confined to the focal volume. Two-photon polymerization is the basis for a liquid-based three-dimensional printing process [7] in which either unpolymerized regions (negative photoresist) or polymerized regions (positive photoresist) are washed away with a developing solvent, leaving only the fabricated structure. This technique has been used to fabricate micro-optic components [8], and hydrogel-based cellular scaffolds [9]. Two-photon crosslinking of proteins in solution has also been reported, including bovine serum albumin [10] and collagen [11]. However, the use of two-photon processes for the precise, internal modification of soft condensed materials, such as biological tissues, has not been reported. A major challenge has been a lack of appropriate tools to monitor and characterize crosslinked regions nondestructively. One study attempting two-photon collagen crosslinking (2P-CXL) in corneal tissue reported changes in collagen second harmonic generation (SHG), but with no evidence of crosslinking [12]. Two other studies have shown the feasibility of 2P-CXL for the stiffening of biomaterials, including artificial cardiac tissue [13] and collagen hydrogels [14]. However, the depth-selective ability of 2P-CXL was not demonstrated, since conventional indentation testing requires physical contact with the sample, limiting the measurements to only surface properties.

In this Letter, we demonstrate spatially selective 2P-CXL of intact tissue for the first time to our knowledge. We use femtosecond laser pulses and riboflavin, a clinically approved photosensitizer, to crosslink a small volume within animal corneas. We characterize collagen crosslinking non-invasively in situ with multiphoton imaging and after the procedure with confocal Brillouin microscopy. Our results demonstrate the ability to alter the microstructure and mechanical modulus locally in the cornea by 2P-CXL with microscopic resolution and illustrate the unique advantage of Brillouin microscopy in evaluating 2P-CXL induced changes nondestructively.

Before characterizing 2P-CXL, we first characterized standard one-photon CXL (1P-CXL) with riboflavin and UV light (365 nm, $3\mathrm{mW}/{\mathrm{cm}}^2$, 30 min). We imaged riboflavin fluorescence and backscattered SHG from collagen fibers in excised, de-epithelialized bovine corneas (see Supplement 1). We used a home-built two-photon microscope with a 1.0 NA water immersion objective lens (Fig. 1). Our light source consisted of a mode-locked Ti:Sapphire laser delivering 810 nm light at 150 fs pulses and an 80 MHz repetition rate.

 

Fig. 1. Home-built two-photon microscope. The cornea sample was compressed with a coverslip. The SHG (blue, 400 to 410 nm) and fluorescence (green, 485 to 555 nm) were collected through the appropriate dichroic mirrors and filters by separate photomultiplier tubes.

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Application of a 0.1% riboflavin, 10% dextran solution to the excised corneas for 30 min caused significant dehydration, resulting in a decrease in corneal thickness from $713\pm 37\mathrm{\mu m}$ before dye application to $473\pm 41\mathrm{\mu m}$ after staining, which was unchanged at $485\pm 19\mathrm{\mu m}$ following CXL (Fig. 2). Fluorescence imaging revealed riboflavin dye penetration over 300 μm below the cornea surface, which was bleached significantly following 1P-CXL. The mean SHG intensity increased following dye application and CXL but was highly variable. The changes in SHG could be attributed to the decreased attenuation by riboflavin absorption, as well as the alteration of collagen fiber distribution, structure and morphology; the latter can significantly affect the intensity and angular distribution of SHG [15]. To explore this further, we next investigated whether the collagen morphology obtained from SHG imaging could be a suitable metric for CXL.

 

Fig. 2. Assessment of 1P-CXL with multiphoton imaging. (a) Reconstructed depth profile of riboflavin-soaked corneas before and after CXL. Green: riboflavin fluorescence, Blue: backscattered collagen SHG. (b) Summary of changes, with mean $\pm$ std. error reported. *, $p<0.05$. (c) Top: distribution of collagen fiber orientations determined by 2D-FFT. Red: fitting to von Mises distribution. Bottom: standard deviation ($\sigma$) of the fiber orientation. A uniform fiber distribution is defined to have a $\sigma$ of 0.5 radians (dotted line). ****, $p<0.0001$.

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The two-dimensional fast Fourier transform (2D-FFT) of the SHG image can be used to quantify the fiber orientation [16,17]. Following filtering to suppress background autofluorescence (Fig. S1), the magnitude of the 2D-FFT gives the integrated power spectrum, which can be interpreted as the number of fibers as a function of the orientation angle [17]. The data was fit to a bimodal von Mises distribution, from which the variance specifies the alignment of the fibers. We found that the orientation distribution of the collagen fibers increases markedly with the application of the riboflavin-dextran solution, such that the fiber orientations are essentially randomly distributed [Fig. 2(c) and Fig. S2]. Subsequent 1P-CXL did not cause any appreciable change to the distribution of the fiber orientations. Our results agree with qualitative observations made previously [18,19], but it is unclear whether any changes in fiber orientation are sustained with corneal rehydration [20].

To determine the parameters needed for 2P-CXL, we first measured the two-photon photobleaching kinetics of riboflavin in solution as a function of the incident laser power from 104 to 230 mW. The power dependence of the bleaching rate was found to be nearly quadratic ($n=1.9\pm 0.3$), indicating a two-photon absorption mechanism [Fig. 3(a)]. Next, we scanned a laser beam at 200 mW over a cross-sectional plane 50 μm below the cornea surface for 20 min (6.6 ms per focal spot) and measured both photobleaching and SHG. We observed a small yet significant $4.3\pm 1.6\%$ decrease in fluorescence ($p<0.01$) [Fig. 3(b)]. Compared to riboflavin bleaching in solution, photobleaching in tissue is much slower, considering the light scattering and rapid diffusion of non-bleached riboflavin ($\sim 40{\mathrm{\mu m}}^2/\mathrm{s}$) [21]. Both the mean SHG intensity and the collagen fiber orientation were unchanged with 2P-CXL [$p>0.05$, Fig. 3(b)]. The insensitivity of SHG to 2P-CXL of a single plane underscores the need for a more specific metric for evaluating crosslinking.

 

Fig. 3. Determination of laser parameters for 2P-CXL. (a) Two-photon bleaching time constant (1/e) in seconds as a function of the incident laser power (mW). Mean $\pm$ std. error are reported. Black line is linear fit to $\ln (\tau )$ vs. ln(power). (b) Changes in SHG and fluorescence with 2P-CXL. Mean (solid line) $\pm$ std. error (dotted line) are reported. Inset: change in the standard deviation ($\sigma$) of fiber orientations. (c) There is no tissue damage with 20 min irradiation at 207 mW, but there is significant photodamage with 5 min irradiation at 312 mW (white arrow). SHG (blue) and fluorescence (green) channels are shown. Scale bar: 50 μm.

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Finally, we compared the laser parameters used for conventional 1P-CXL with 2P-CXL, assuming that an equal number of molecular excitations is needed to elicit the same crosslinking effect (see Supplement 1). We found that at an average power of 100 mW, 2P-CXL of our $(290\times 480){\mathrm{\mu m}}^2$ scanning area would require $\sim 2.1\times {10}^4\mathrm{J}/{\mathrm{cm}}^2$. In subsequent samples characterized with Brillouin microscopy, 2P-CXL was performed on riboflavin-stained corneas at 104 mW with scanning times of 10 min per plane ($\sim 4.3\times {10}^4\mathrm{J}/{\mathrm{cm}}^2$), accounting for losses due to optical aberration and light scattering. While no discernable tissue damage was observed at 207 mW for up to 20 min, irradiation with 312 mW for 5 min resulted in tissue damage, likely due to nonlinear photochemical and photothermal effects causing low-density plasma formation [22] [Fig. 3(c)].

Unlike measurements of riboflavin fluorescence and collagen SHG, Brillouin microscopy contrast is directly dependent on the local hydromechanical properties of the tissue. Brillouin microscopy has been used to characterize normal and keratoconic corneas in vivo [23], as well as the effects of 1P-CXL ex vivo [24]. This technique relies on the scattering of photons by spontaneous acoustic waves, which are sensitive to changes in the compressibility of the local microenvironment, modulus of collagen fibers, and the crosslink density of the network [Fig. 4(a)]. The resulting shift in optical frequency is detected using a high-resolution spectrometer. Using a previously described confocal Brillouin microscope [25] [Fig. 4(b)], we imaged a cornea immediately following 2P-CXL-induced stiffening of an arbitrary three-dimensional region centered around 50 μm below the surface (10 successive planes, 3 μm apart). The distribution of riboflavin over these depths is nearly uniform (Fig. S3).

 

Fig. 4. Brillouin microscopy. (a) Brillouin scattering is due to interaction of photons and spontaneous acoustic waves. For the 2P-CXL region, the acoustic velocity is higher for a given phase-matched phonon wavelength. (b) Confocal Brillouin microscope connected to a two-stage, apodized, virtually imaged phased array spectrometer.

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Scanning large areas of the cornea, we readily identified a region with significantly increased Brillouin frequency shifts, indicating local stiffening (Fig. S4). Focusing on the stiffened region, we then took $x–z$ and $x–y$ scans, which were reconstructed as a three-dimensional volume (Fig. 5 and Visualization 1). A rectangular volume spanning approximately $(205\times 325\times 50){\mathrm{\mu m}}^3$ was found to be significantly stiffer than the surrounding regions [Fig. 5(c)]. These dimensions compare to a volume of $(290\times 480\times 30){\mathrm{\mu m}}^3$ that was irradiated with our two-photon microscope. While the total volume of the 2P-CXL region is nearly conserved, the discrepancy in area and thickness was attributed to flattening of the sample with a coverslip during 2P-CXL, which was required for use with our high NA, water immersion objective. In contrast, with Brillouin microscopy, the cornea was imaged in its natural condition without artifacts caused by coverslip compression. Distortion from a rectangular shape can be attributed to optical aberrations leading to a reduced intensity at the periphery of the scanning area in the two-photon microscope.

 

Fig. 5. Brillouin microscopy of a 2P-CXL-treated cornea. (a) $x–y$ images of the 2P-CXL region, at 15, 45, 75, and 105 μm below the cornea surface. Scale bar: 50 μm. b) $x–z$ image, where top refers to the surface of the cornea. (c) Three-dimensional view of the 2P-CXL region reconstructed and interpolated from thirteen $x–y$ images (selected images shown in a) at 10 μm intervals in $z$.

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The mean Brillouin shift over the anterior 300 μm of the untreated cornea was $8.15\pm 0.04\mathrm{GHz}$, in agreement with previous ex vivo studies measured at the same wavelength of 532 nm [24,26]. The depth profile of the 2P-CXL region and the surrounding non-irradiated region is plotted in Fig. 6(a). Over depths of 30 to 80 μm below the surface, 2P-CXL increased Brillouin frequency shifts by 0.3 to 0.6 GHz. This corresponds to an estimated increase of 0.2 to 0.4 GPa in the longitudinal modulus (see Supplement 1). The relative Brillouin shift across the irradiated volume was fitted to a Gaussian function, giving a peak of $548\pm 7\mathrm{MHz}$ at a depth of $55.7\pm 0.3\mathrm{\mu m}$ and an FWHM of $48.7\pm 0.7\mathrm{\mu m}$ [Fig. 6(b)]. This gradual change in the Brillouin shift is largely due to the limited axial resolution of the 0.3-NA objective lens used. The degree of stiffening achieved by 2P-CXL is comparable to a $\sim 0.5\mathrm{GHz}$ increase previously measured in the anterior cornea with conventional 1P-CXL [24]. Notably, 1P-CXL causes significant stiffening over nearly the entire stromal depth over $>500\mathrm{\mu m}$ [Fig. 1(a)], while 2P-CXL-induced stiffening can be localized to target depths [Fig. 5(b)].

 

Fig. 6. Quantification of 2P-CXL-induced stiffening. (a) Mean ($\pm$ st. dev.) Brillouin shift of the crosslinked and non-irradiated regions of the cornea as a function of depth. (b) Relative Brillouin shift obtained from subtracting the profiles in (a). Fit to Gaussian function ($R^2=0.99$).

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In summary, we have demonstrated three-dimensional selective crosslinking in corneal tissue. Our technique involves (1) in situ 2P-CXL of an arbitrary three-dimensional structure, and (2) verification and characterization with confocal Brillouin microscopy. Compared to conventional procedures for evaluating 2P-CXL-induced stiffening, Brillouin microscopy does not require perturbation of the surrounding non-crosslinked regions. Other non-invasive approaches for measuring tissue elasticity include elastographic techniques based on ultrasound [27] or magnetic resonance imaging [28], but at typical spatial resolutions of hundreds of microns to several millimeters, respectively. An emerging technique is optical coherence elastography, which can be used to extract viscoelastic properties by optical measurement of mechanically induced tissue deformations [29].

Our current technique is limited by the long irradiation times needed for 2P-CXL, about 1 h per ${\mathrm{mm}}^2$. Faster 2P-CXL could be achieved by using wide-field two-photon microscopy techniques such as light sheet or Bessel beam plane illumination [30,31]. Although riboflavin was chosen for direct comparison with conventional corneal CXL, it has a relatively low two-photon absorption cross-section at 810 nm ($<1\mathrm{GM}$), which can be enhanced by a factor of $\sim 2–3$ by excitation at 760 nm [14,32]. More efficient two-photon photosensitizers are also in development [33]. Increasing the peak irradiance by the use of shorter pulse width or lower repetition rate lasers could shorten 2P-CXL times, but careful selection of laser parameters is needed to avoid nonlinear photodamage [Fig. 3(c)]. Another challenge is the acquisition time needed for Brillouin microscopy, typically over 30 min per image in this work. The exposure time could be decreased with a better signal-to-noise ratio, such as by improving the filtering of background Rayleigh scattering [34]. Future development of a combined two-photon and Brillouin imaging setup would enable the real-time measurement and feedback control of 2P-CXL-induced stiffening.

Our current 2P-CXL technique can be employed for applications involving micron-scale target regions, such as microfabrication and patterning, as well as precise laser surgeries. Nonlinear excitation permits three-dimensional control of crosslinking, while the use of NIR light enables deep tissue penetration. 2P-CXL can be readily applied in ophthalmology, such as for treatment of thin ectatic corneas ($<400\mathrm{\mu m}$) [6], corneal flap bonding post-LASIK surgery, and the selective modulation of corneal curvature for refractive error correction.