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.