The refractive index variation in tissues, quantified by the ratio m = njn0, determines light scattering efficiency. For example, in a simple tissue model, such as dielectric spheres of equal diameter 2a, the reduced scattering coefficient is described by [40]:
m’s = ms (1- g) = 3.28 na2 p(2na/A)0,37 (m – 1)209, (3.14)
where p is the volume density of the scatterers, g is the scattering anisotropy factor, and l is the light wavelength in the scattering medium. At equalizing (matching) of refractive indices (RI) of the interstitial fluid, n0, and scatterers, ns, m ^ 1, p’s ~ps ^ 0. Skin transmittance, T = I(d)/I0 ~ exp{-[3дl(д,, + ^a)]1/2d} [Eq. (3.7)], can be increased substantially by reduction of the scattering coefficient, because for a native skin in the visible/NIR range ц’ >> цл. One of the ways to decrease the scattering coefficient of skin is by impregnating it with a solution [optical clearing agent (OCA)] with RI, nOCA, higher than n0 [1,41]. If a hyperosmotic OCA is applied topically to the skin, besides its diffusion into skin, tissue water will flow outside from a tissue and corresponding tissue dehydration will take place, this action will lead to additional matching of RI of scatterers relative to the background and more effective packing of scatters (tissue shrinkage). Both processes provide more effective light transport through skin.
The excellent diffusional resistance of the skin stratum corneum (SC) makes the transdermal delivery of immersion agents and water lost by skin difficult [42]. The diffusion of water across the SC is a passive process that can be modified during the application of hyperosmotic OCAs. The water content of the innermost layer of the SC is in equilibrium with the adjacent moist granular layer. The outside cell layer, however, is in equilibrium with the environment, and it is certainly drier than the innermost cornified layer. Dermis is the thicker layer of the skin, which is mostly fibrous tissue well-supplied by blood, and thus can be easily impregnated by exogenous or endogenous liquids (immersion agents). Subcutaneous tissue contains a big portion of fat cellular layer, which is much less penetrative for diffusing molecules than dermis. Such specific structure of skin defines the methodology of its effective optical clearing, which is related to the immersion of refractive indices of scatterers (keratinocytes components in epidermis, collagen and elastin fibers in dermis) and ground matter [41].
Experimental studies of optical clearing of skin using glycerol, glycerol-water solutions, glucose, propylene glycol, polyethylene glycol, DMSO, sunscreen creams, cosmetic lotions, gels, and pharmaceutical products were recently overviewed [1,41,43]. In vivo topical application of these agents made human skin more optically translucent within a time period, from a few minutes to a few hours.
To enhance OCA permeation through SC, a number of specific physical procedures, such as heating, electrophoresis, sonophoresis, and laser-induced stress, as well as chemical enhancers, such as oleic acid and DMSO, are usually applied. A method of accelerating penetration of the index-matching compounds by enhancing skin permeability through creating a lattice of micro-zones (islets) of limited thermal damage in the SC was recently proposed [44]. A combination of a flashlamp system (EsteLux, Palomar Medical Technologies, Inc.) and a specially designed applique with a pattern of absorbing centers (center size ~75 ^m, lattice pitch ~450 ^m) has been used to create the lattice of islets of damage (LID). Several index-matching agents, including glucose and glycerol, have been tested. A high degree of optical clearance of a full-thickness pig, rat, chicken, and human skin in vitro and in vivo has been demonstrated with 40% glucose and 88% glycerol solution after creating a LID with a few optical pulses (fluence 14-36 J/cm2, 20 ms pulse duration).
To provide faster and more effective skin optical clearing, an intradermal injection can be applied. Figure 3.14 shows the reflectance spectra and the corresponding time-dependent reflectance for a few spectral components measured for a human healthy volunteer at intradermal injection of 40% glucose solution [45]. The reflectance spectra are determined by the diffusion reflection of the skin layers with the well-pronounced bands caused by blood absorption. Within one hour after glucose injection, the skin reflection coefficient decreases in average by a factor of 3.8 and then exhibits a slow increase, which indicates that glucose is eliminated from the observation area, and the skin reflectance tends to restore
itself to the initial level. Basing on this results and skin model, it was suggested that the main contribution to clearing in the first stage (first hour) is due to the RI matching between collagen fibrils of the dermis (n = 1.46) and the interstitial space (initially n = 1.36) to which glucose (n = 1.39) diffuses.
For applications, it is important that skin preserves transparency (low reflectance) for a few hours after injection, which is defined by predominant diffusion of glucose along the skin surface, because the upper and lower layers of the skin—epidermis and fat—have much lower (a few orders) permeability for glucose than dermis. It is seen from Fig. 3.14 that at dermal clearing the contrast of hemoglobin absorption bands is significantly higher than for normal skin, but for prolonged immersion (curve 3) the contrast is again not very high. This is important for the optimization of clearing time at imaging of tissue abnormalities associated with hemoglobin or other absorbers.
Because of the limitation of probing the depth of OCT imaging (1-2 mm for skin), its combination with OCA immersion can be a useful technology for skin diagnosis and monitoring. This is illustrated by the OCT images of human skin with psoriatic erythrodermia acquired before, and in some time after application of glycerol (Fig. 3.15) [46]. In one hour
of glycerol application, OCT image differs from the initial image in greater penetration depth and better contrast. These image improvements facilitate identifying of important morphological phenomenon of acanthosis.
Squeezing (compressing) or stretching of skin produces a significant increase in its optical transmission [1,41,43]. The major reasons for that are the following: (1) increased optical tissue homogeneity due to removal of blood and interstitial fluid from the compressed site; (2) more close packing of tissue components causes less scattering due to cooperative (interference) effects; and (3) less tissue thickness.
Spectral properties of skin can be effectively controlled by applying an external localized pressure when UV induced erythema (skin redness) is developed [7]. The intensity of skin reflectance and autofluorescence is well-controlled at pressure applied to the skin site. Due to more effective fluorescence attenuation by blood hemoglobin at more intensive erythema, skin compression more effectively increases fluorescence output.
The light propagation in human skin at mechanical tension was studied in vivo using diffuse reflectometry [22]. It was found that intact skin has its own anisotropy which is believed to be caused by the preferential orientation of collagen fibers in the dermis, as described by Langer’s skin tension lines, and at skin external stretching, scattering coefficient and corresponding light back-reflectance and transmittance can be effectively controlled. At external forced tension, more significant damping of scattering along the direction of mechanical stress was determined. The reduced scattering coefficient varied by up a factor of two between different directions of light propagation at the same position.
The measurements of the deformations and applied loads and estimating the biomechanical properties of tissue are critical to many areas of the health sciences, including monitoring of the tension in wound closures, skin flaps, and tissue expanders [47]. Such measurements which can be provided by detection of the polarized light reflectivity will allow surgeons to treat wounds more successfully by minimizing scar tissue and maximizing the speed of treatment, by letting them know how much the skin can be stretched at each treatment step. In vivo human experiments showed that the specular reflection from skin changes with stretch [47]. For small values of stretch, the specular reflectivity measured for He-Ne laser (l = 633 nm) beam with the 45° angle of incidence increases linearly with strain. The linear relationship between applied stretch and polarized reflectivity can be understood if the skin surface is approximated by a sinusoidal profile in the resting stage. Stretching reduces amplitude and increases spatial scale of skin profile, thereby making it smoother and flatter, resulting in a corresponding increase of reflectivity. For larger stretches [for strains above 8.8% (5-mm stretch)] for the human subject tested, the dependence is saturated and even goes down. The stretches in two perpendicular directions (parallel and perpendicular to the long axis of the forearm) yield good correlation between stretch and reflected light intensity and shows that skin has anisotropic properties, which can be detected by light reflection [47].
A reproducible effect of temperature between 25 and 40°C on the reduced scattering coefficient of human dermis and subdermis was found in ex vivo study in the NIR [18]. For dermis, the relative change in the reduced scattering coefficient showed an increase [(4.7 ± 0.5) x 10-3°C-1] and for subdermis a decrease [(-1.4 ± 0.28) x 10-3°C-1]. It was hypothesized that the observed positive and negative temperature coefficients of scattering for dermis and subdermis are connected with differences in their structural components.