1 INTRODUCTION
Human skin is an inhomogeneous organ comprised of three main layers: epidermis, dermis and adipose tissue. Each of these layers differs in thickness and morphology. Playing a role of an interface between the body and the environment, the skin performs physiological barrier function, protecting the body from ultraviolet rays, external physical impact and retaining water inside the body.[1]Since the skin is actively exposed to external irritants, it is susceptible to various pathological processes that can affect different layers of the skin, depending on the type and stage of the lesion. Due to the importance of the skin for the proper functioning of the body, the non-invasive diagnosis of skin diseases is an extremely important task, to which many studies are devoted. Optical methods based on the interaction of light with biological tissues have become a good potential addition or replacement to existing invasive methods (such as histopathological studies) due to their non-invasiveness, high sensitivity, the ability to obtain information in real time (without time consuming preparation of histological samples), as well as potential for clinical implementation.[2, 3]
Line-field confocal optical coherence tomography (LC-OCT) is a recently developed optical imaging method that combines the advantages of OCT and confocal microscopy, providing three-dimensional (3D) images of tissue with a quasi-isotropic resolution of ~1 µm that is high enough to distinguish the cellular skin structure.[4,5] The applicability of LC-OCT to the diagnosis of skin diseases (including skin cancers) has been widely demonstrated.[6-8] Besides information on tissue morphology, the LC-OCT images also contain information on tissue optical properties. Since LC-OCT is an interferometric method, the image is obtained by registering low-coherence light backscattered or reflected by the tissue. This signal depends on three main optical properties of the imaged tissue: absorption µa(λ) and scattering µs(λ) coefficients, which determine the fraction of light scattered or absorbed over a unit path length, respectively, and the scattering anisotropy factor g(λ) that is the mean cosine of scattering angles along the photon trajectory. The in-depth LC-OCT image intensity \(I\left(z\right)\) depends on these optical properties and usually can be described (although only partially since the scattering anisotropy, for example, is not taken into account) by a Beer-Lambert law in the single-scattering regime:
\(I\left(z\right)\propto\ \exp^{-2µ_{t}(\lambda)z}\), (1)
where z is depth in mm and µt(λ) = µs(λ) + µa(λ) is the total attenuation coefficient of the medium in mm-1.[9] Those properties are determined by the size, density and shape of the tissue constituents (cells, collagen and elastin fibers, etc.). Assessing the optical properties of skin is critical for tissue characterization and quantification of structural changes associated with the pathological process.
The method most widely reported in the literature for extractingµt(λ) with conventional OCT techniques is to fit an exponential decay curve to the depth-dependent average intensity profile.[10] However, only few works were dedicated to separate assessment of scattering, absorption and scattering anisotropy from OCT images.[11] Such an assessment may allow one to obtain more comprehensive quantitative information about structural changes induced by a particular process in the skin than when estimating the integral µt(λ) coefficient. But this approach becomes complicated to implement when it comes to multilayered samples due to conventional OCT features and the concept of backscattered light as a fixed fraction of the attenuated light (which is assumed when extracting µt(λ) coefficient).[12] In contrast, a model based on Monte-Carlo simulations, developed by Jacques et al .[13] allows for a simple extraction ofµs(λ) and g(λ) from focus-tracking OCT techniques and confocal microscopy. This model was validated on phantoms with pre-defined optical properties[14] and later applied to skin.[15] Since LC-OCT is a combination of OCT and confocal microscopy techniques, this technique is suitable for application of the aforementioned model. In her work, Waszczuket al .[16] demonstrated that, with preliminary calibration using a phantom with known optical properties, it is possible to extract skin µs(λ) andg(λ) optical properties from 3D LC-OCT images of monolayered and bilayered phantoms.
However, the use of OCT (as well as other optical methods) for skin diagnosis is limited by strong light scattering in biological tissues, e.g., in skin. Scattering reduces the image contrast and resolution, lowering the possible diagnostic potential of such methods. This scattering originates from the inhomogeneities of skin layers (intralayer and interlayer inhomogeneities), leading to a mismatch in refractive indices (RI) between the tissue constituents and the interstitial fluid. To overcome this limitation, a tissue optical clearing (OC) method was proposed, based on the use of osmotic chemicals, called Optical Clearing Agents (OCA), whose RI was close to that of tissue solid material.[17,18] Being usually topically applied (but also can be injected into the tissue), OCA cause skin dehydration, followed by replacement the interstitial fluid with OCA and reversible collagen dissociation. It results in reduced scattering of treated tissue, leading to increased imaging depth and contrast.[19]
Translation of skin OC into clinical use, however, is connected with the need to comply with established regulations on the use of drugs, especially if the possible application will be performed on a lesional tissue. Since at pure concentrations OCA have been reported to have undesired side effects in vivo ,[20,21]their concentrations must be reduced in order to pass the threshold for clinical admission and biocompatibility.[22] But the low concentration of OCA does not allow to reach a sufficient clearing effect. To compensate for lower OCA concentrations in biocompatible applications, they can be used in conjunction with so-called chemical permeation enhancers (CPE), which are chemicals capable of temporarily disrupting skin barrier functions. The chemicals most commonly used as CPEs are alcohols,[23]dimethyl sulfoxide and fatty acids (Oleic acid).[24] There is also a large number of physical methods to enhance the skin permeability for OCA, such as microdermabrasion and therapeutic ultrasound,[25]which can be combined with CPE for a more efficient biocompatible effect of OCA.
Such possibility of in vivo clearing of human skin using biocompatible OCA concentrations has been addressed in our previous study,[26] where nine mixtures of one of three OCA compounds (Polyethylene Glycol (PEG), Sucrose and Glucose water solutions) in combination with one of three CPE compounds (Propylene Glycol (PG), Dimethyl sulfoxide (DMSO) and Oleic acid (OA)) were used with skin microdermabrasion and ultrasound to test their effectiveness in increasing LC-OCT in-depth image intensity and contrast. It was demonstrated that all tested OCA compositions caused an increase in the ratio between mean intensity and contrast extracted from 3D LC-OCT images. It was assumed that the main reason for such effects was a change in skin optical properties, i.e., a decrease of the scattering coefficient, caused by OCA. However, that assumption was not experimentally confirmed.
Taking into account the method developed by Jacques at al .[13], and later validated for the LC-OCT technique by Waszczuk et al .,[16] the goal of the current study was to quantify, using LC-OCT, the values of scattering coefficient µs(λ) and scattering anisotropy factor g(λ) and their modifications caused by in vivo biocompatible optical clearing of human skin.
2 MATERIALS AND METHODS
In this study, the composition of OCA compounds, the skin sites under investigation and the imaging modality was similar to our previously published work, where some complementary information could be found.[26]
2.1 Optical Clearing Agents and permeability enhancers
The choice of chemicals for this study was based on literature data on OCA currently used for human skin OC experiments.[17] Three chemicals from alcohol and sugar groups were used as OCA compounds: Polyethylene Glycol - 400 (PEG, Sigma-Aldrich, USA), 3M aqueous solutions of Sucrose (Sigma-Aldrich, USA) and Glucose (Sigma-Aldrich, USA). In order to satisfy the possible clinical admission requirements, OCA compounds did not exceed threshold of a concentration for topical application in the form of a solution, established by FDA. Thus, the data from FDA inactive ingredients database was used to fix the concentration of OCA mixture compounds, mentioned hereafter.[22] As there was no information about maximum concentration for glucose and sucrose solutions, a value (v/v) of 50% was established for the current study as it was previously reported as the most efficient one for OCT-assessed optical clearing.[27] To increase the in vivo efficiency of reduced concentrations of OCA compounds, they were mixed with three compounds with permeation-enhancing properties (CPE) from various chemical groups, such as alcohols, organic solvents and fatty acids: Propylene Glycol (PG, Sigma-Aldrich, USA), Dimethyl sulfoxide (DMSO, Sigma-Aldrich, USA) and Oleic acid (OA, Sigma-Aldrich, USA). Nine resulting mixtures of OCA and CPE and their corresponding compound concentrations are presented in Table 1 . If it was not possible to mix OCA and CPE compounds only without exceeding FDA concentration thresholds, either complementary amount of distilled water or second CPE (namely PG) was added to the mixture. Additional information (such as the concentration threshold of each chemical used) can be found in our previous study.[26]
TABLE 1 . Nine mixtures of OCA and CPE with corresponding compounds concentrations (%, v/v) that meet the FDA-allowed concentration threshold.