Estimation of the Optimal Wavelengths for Laser-Induced Wound Healing

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1 Lasers in Surgery and Medicine 42: (2010) Estimation of the Optimal Wavelengths for Laser-Induced Wound Healing Rinat Ankri, 1,2 Rachel Lubart, 1,2 * and Haim Taitelbaum 2 1 Department of Chemistry, Bar-Ilan University, Ramat-Gan, Israel 2 Department of Physics, Bar-Ilan University, Ramat-Gan, Israel Background and objectives: According to earlier in vitro low level laser therapy (LLLT) studies, wavelengths in the red and near infrared range, that are absorbed by cytochrome oxidase, stimulate cell growth and hence wound healing. Wavelengths in the blue region that are absorbed by flavins were found to exert a bactericidal effect that is very important for treating infected wounds. However, as far as therapeutic application of light is concerned, penetration into the tissue must be considered. For this purpose we estimated the penetration depth as a function of the relevant wavelengths, using the formulae of the photon migration model for skin tissue. Methods: We use the photon diffusion model, which is an analytical model for describing light transfer in biological tissues. We refer to the most common chromophores in human tissue and evaluate their volume fraction and concentration in skin cells. These empirically estimated mean wavelength-dependent absorption coefficients are then substituted in the theoretical expressions for the optical penetration depth in the tissue. The wavelengths, for which the penetration depth is the highest, are the optimal wavelengths to be used in wound healing treatments. Results: Our model suggests that the optimal wavelengths for therapeutic treatments are in the red region with a local maximum at 730 nm. As to the blue region, a local maximum at 480 nm was found. Conclusion: Light at 480 nm should be used for treating infected wounds followed by 730 nm light for enhancing wound closure. Lasers Surg. Med. 42: , ß 2010 Wiley-Liss, Inc. Key words: wound healing; skin tissue; photon migration model; light penetration depth INTRODUCTION Wavelengths in the visible and NIR light are used in the therapeutic field, such as photo dynamic therapy (PDT), that uses light to damage tumor cells, and low level laser therapy (LLLT), that uses visible-nir light to biostimulate the cell [1 4]. In photobiostimulation, the interaction of light with the cell is ascribed to the excitation of intracellular chromophores like endogenous porphyrins, mitochondrial and membranal cytochromes and flavoproteins. These chromophores transfer their excited electrons to nearby O 2, thus generating low amounts of reactive oxygen species [5], which play an important role in the activation and control of many cellular processes, such as the release of transcription factors, gene expression, muscle contraction, and cell growth [6,7]. Recently, it has been shown that wavelengths in the blue region ( nm), induce generation of higher amounts of ROS than wavelengths in the red one ( nm) [6]. This phenomenon is exploited for sterilization purposes, such as bacterial eradication in contaminated wounds [7 11]. Moreover, very recently blue light has been demonstrated to be mostly responsible for NO formation by endothelial cells [12]. Since NO formation leads to vasodilatation and subsequent increase in microcirculatory blood flow, the use of blue light might be of great importance for wound healing of diabetic and venous ulcers. However, as far as therapeutic application of light is concerned, penetration inside the tissue must be considered. In the present work we do so by using the photon migration model [13]. We apply this model to the dermis layer which is the exposed layer in the wounded tissue, neglecting the upper epidermis layer. The mean absorption coefficients of the dermis layer are calculated and substituted in the theoretical expressions for the penetration depth, thus getting estimation for the penetration depth as a function of the visible wavelengths. The optimal wavelengths for clinical treatments are those that penetrate deepest inside the tissue, within the relevant wavelength range. There are several methods and numerical techniques to consider light propagation in tissues, such as Monte Carlo simulations, finite differences approaches, diffusion approximation of radiative transfer equations, etc. (see, e.g., Tuchin s book [14]). The main advantage of the random-walk on the lattice model, over other methods, is the ability to derive analytical expressions for various quantities of interest, in particular those related to the important issue of the light penetration depth inside the tissue [13,15,16]. *Correspondence to: Rachel Lubart, Department of Chemistry, Bar-Ilan University, Ramat-Gan 52900, Israel. lubartr@mail.biu.ac.il Accepted 17 June 2010 Published online 15 September 2010 in Wiley Online Library (wileyonlinelibrary.com). DOI /lsm ß 2010 Wiley-Liss, Inc.

2 LASER-INDUCED WOUND HEALING 761 METHODS Photon Migration in a One-Layer Biological Tissue In the photon migration model, the tissue is represented by a semi-infinite cubic lattice with two adjustable parameters, namely its scattering and absorption factors. The simplest version of the model refers to a one-layer tissue and was successfully used to predict the intensity of re-emitted flux, the mean trajectory length of detected photons and the penetration depth of photons inside the tissue [15]. Figure 1 shows the one-layer version of the photon migration model: The tissue is modeled as a three dimensioned discrete lattice with transverse coordinate, r ¼ (x,y), and a positive z-axis pointing into the tissue. Photons are injected at the origin, then diffuse randomly within the tissue, eventually either reach the surface z ¼ 0, where they can be detected, therefore disappearing from the system, or they are being absorbed inside the tissue. The scattering of the photon in the lattice is assumed to be isotropic and is described by a random change in the photon s direction of motion on the cubic lattice. The absorption, occurring between nodes of the lattice, is described by Beer s law, so that the survival probability of a given photon is exp ( m) per step on the lattice. In order to correlate the absorption coefficient used in the model, m, which is per step on the lattice, and the conventional absorption coefficient used in literature, a, which has conventional inverse length units, one should multiply m with the lattice unit length. In the NIR region, it was found that one lattice unit length is equivalent to a length of about mm on the biological tissue. For LLLT treatments on wounded skin tissues we have chosen a length of 0.1 mm for one lattice step. Among other statistical characteristics presented in the model, the average depth reached by photons that are eventually being absorbed inside the tissue is hzi a ¼ p 1 ffiffiffiffiffi ð1þ 6m where m is the model parameter for the wavelengthdependent mean absorption coefficient of the tissue and 1/6 is the lattice diffusion constant, where the randomwalker has six possible directions of motion from each lattice site. This average depth has units of lattice length, since the diffusion constant units are (length) 2 per lattice step, and m is the absorption per step. hzi a is translated to conventional units by multiplying m with the lattice constant length; 0.1 mm. Model Application to the Human Skin In order to calculate light penetration depth for wound healing purposes, we apply the one-layer photon migration model to the dermis layer of the skin, since it is the exposed layer in wounded skin. Indeed, the dermis is an inhomogeneous layer since it is consisted from two main sub-layers; the papillary dermis, which is the upper layer in the dermis, and the reticular dermis. Still, the dermis layer can be considered homogeneous relating to the fibroblast cells and the blood vessels, which contains the main chromophores in the visible region and that are homogeneously distributed in both layers [17]. To apply the model to the dermis skin layer, the necessary first step is to estimate the absorption coefficients of each of the elements that constitute skin tissue cells. Specifically, we refer to their most absorbing chromophores in the visible region that have a significant role in ROS formation by illuminated cells, as mentioned in the introduction. Those chromophores are hemoglobin (Hb), cytochrome c (Cyt c), and riboflavin (RF). Each of these chromophores is characterized by an absorption coefficient (i.e., wavelength dependent) and has volume fraction (i.e., percentage in skin tissue) describing its appearance in the dermis layer. Those chromophores are also characterized by their concentration in a skin cell, as explained in the Volume Fraction and Cell Concentration of Skin Chromophores Section. H 2 O molecules were also added to the mean absorption coefficient calculation since they own a significant absorption in the NIR region ( nm), therefore, in the next paragraphs, H 2 O molecules will be referred as chromophores. In order to obtain mean absorption coefficients for the dermis we average the absorption coefficients of the above-mentioned chromophores. The wavelength-dependent absorption coefficient values are multiplied by their volume fraction and cell concentration in skin tissue. In the next two paragraphs we describe the way the absorption coefficients, volume fraction, and cell concentration data are obtained for each of those chromophores. RESULTS Chromophore s Absorption Coefficients Fig. 1. Discrete lattice representing the biological tissues in the photon migration model and typical photon trajectories in a one-layer tissue. The absorption coefficients for Hb and H 2 O were adopted from Oregon Medical Laser Center website [18]. The absorption coefficients of Cyt c and RF were experimentally calculated: solutions of Cyt c molecules, with known concentration of 10 mm, and solutions of RF molecules, with known concentration of 75.3 mm, were spectrally

3 762 ANKRI ET AL. measured using spectrophotometer (SpectrafluorPlus, Tecan) in the visible light range ( nm). The absorption values were translated into absorption coefficients wavelength-dependent values using Beer Lambert linear absorption law: A ¼ LCa ð2þ where A is the absorption values (OD), L the test tube s length (1 cm), C the chromophore s known concentration (M), and a the chromophore s absorption coefficient (1/M cm). The absorption coefficient spectra for both chromophores, RF and Cyt c, are presented in Figure 2a,b, respectively. The next step was to weight those wavelength-dependent absorption coefficients according to their volume fraction (and cell concentration for Hb, Cyt c, and RF) as appear in skin tissue. Volume Fraction and Cell Concentration of Skin Chromophores We first referred to the hemoglobin chromophore. The volume fraction of blood vessels in the dermis layer is 2% [17]. A red blood cell contains 95% Hb, therefore having a strong absorption in the visible light region. The molar concentration of Hb in a blood cell is 2.27E 3 M, based on its molecular weight 66,500 g/mol and its concentration in blood, 150 g/l [17]. Other skin chromophores, as mentioned above, are Cyt c, RF, and H 2 O molecules. Our assumption is that those molecules are contained in all skin cells that are not blood vessels in the dermis. Therefore, in the dermis layer, that contains 2% of blood vessels, the left 98% consist of 70% H 2 O molecules [19] and 30% dry mass of skin cells. Hence, one has 68.6% of H 2 O molecules and 29.4% of dry mass of skin cells in the dermis. As mentioned above, those volume fractions of Cyt c and RF should be multiplied by their concentration in a skin cell. The cellular concentration of RF, 5E 5 M, was taken from literature [20], while the cellular concentration of Cyt c was experimentally determined as follows: fibroblasts (NIH/3T3 cells), which are cells having physical properties similar to skin cells, were grown as described elsewhere [21]. The cytoplasm fluid was isolated from 5.4E7 cells using fibroblasts fractionation as followed: Aragon gas was diffused to the cell solution in order to keep the reduced form of Cyt c molecules. The cells were sediment using a 1,500 rpm centrifuge. The sedimentation was deposited in dry ice and on a hot plate, alternately, and then went through a 20,000 rpm centrifugation for 1 hour. The superior fluid, which is the cytoplasm, was spectrally measured (in the visible light region) to give the absorption spectrum, as presented in Figure 3. One can notice the absorption peaks of the cytosol solution around wavelengths 410 and 550 nm, indicating the presence of Cyt c molecule in fibroblast cytosol solution [22]. From those spectral data and by using Equation (2) we obtained Cyt c concentration in fibroblasts solution whereas:a is the Cyt c s absorption at 410 nm, standing for OD, L the test tube s length, 1 cm, and a the Cyt c absorption coefficient at 410 nm; a (410 nm) ¼ 5.302E4 (1/M cm). Hence, we obtain C(l) ¼ A/(La (410 nm) ) ¼ 9.2E 7 M, while C(l) is the Cyt c concentration in the entire fibroblasts Fig. 2. Absorption coefficient spectra for (a) riboflavin (RF) (b) cytochrome c (Cyt c). Fig. 3. Fibroblasts cytosol absorption spectra (arbitrary units) in the visible region as measured by a spectrophotometer.

4 LASER-INDUCED WOUND HEALING 763 TABLE 1. Main Chromophores in the Dermis Layer, Their Volume Fraction and Concentration in Cell Chromophore Volume fraction in dermis (%) Concentration in cell (M) Hb E 3 Cyt c E 14 RF 5E 5 H 2 O 68.6 solution. The Cyt c concentration in a single fibroblast cell was calculated by dividing C(l) with the number of cells C ¼ CðlÞ=ð5:4e7 cellsþ ¼1:703e 14 M=cell The results from previous paragraphs are summarized and presented in Table 1. One can see that the skin cellular concentration of Cyt c is very small compared to Hb and RF skin cell concentrations, thus it can be neglected in the entire calculation. In addition, our results indicate that RF absorption coefficients are very small, so they can also be ignored. The absorption coefficients of H 2 O are not high in the visible region, but in the NIR region they become significant. Moreover, their volume fraction is very high, so they must be taken into account in the general calculation. In conclusion, our results indicate that the absorption coefficients of Hb were the most dominant factors in the calculation. DISCUSSION AND SUMMARY In the present study, we characterize the top layer in the wounded skin tissue (dermis) by referring to its chromophores absorption coefficients. We obtain the wavelengthdependent absorption coefficients, m(l), of the dermis and use the relevant expressions in the statistical photon migration model to calculate the penetration depth of the visible-nir light in homogeneous skin tissue. The penetration depth graph consists of two different local maxima; one around 480 nm and the other around 730 nm. As mentioned in the Introduction Section, low level red and NIR light, when irradiated on biological cells, stimulate low ROS fluxes which play an important role in the activation and control of many cellular processes, such as the release of transcription factors, gene expression, cell growth [23 25], recovery of damaged nerve cells [26], and more. As the clinical use of LLLT is mostly for wound healing [27] it turns out that 730 nm is the optimal wavelength since it penetrates best the wound. Blue light, which was found to be more effective than red light for ROS Mean Absorption Coefficient and Light Penetration Depth in the Dermal Layer We now apply the above results to the simplest version of the photon migration model. Following Table 1, the mean absorption coefficient of the dermis layer is assumed to follow the weighted form: mðlþ ¼a Hb ðlþ2%ð2:27e 3Þþa H2OðlÞ68:6% þ less absorbing elements ð3þ where a Hb (l) and a H2O ðlþ are taken from the OMLC website [18] and the less absorbing elements refer to Cyt c and RF molecules, which were argued above not to contribute to the skin tissue total absorption. From Equation (3) an absorption coefficient wavelengthdependent graph is obtained and is presented in Figure 4a. The graph is very similar to that of hemoglobin s absorption coefficient graph [18] indicating that the main absorbing chromophore in the dermis layer is the hemoglobin. Assuming that the dermis layer is a homogeneous tissue, we next use Equation (1) to calculate the averaged penetration depth of the irradiated light as a function of wavelength in the visible-nir region. Following the above mentioned relation between a and m (see the Photon Migration in a One-Layer Biological Tissue Section) we multiply the empirical absorption coefficients (a) by 0.1 mm and thus define the modelistic absorption coefficients (m). Figure 4b presents two local maxima at 480 and 730 nm. Fig. 4. Homogeneous dermis tissue. a: The mean absorption coefficient, calculated using Equation (3). b: Penetration depth of irradiated light, calculated using Equation (1).

5 764 ANKRI ET AL. formation [28] and hence for bacteria killing [29], has not yet been clinically used because of its low penetration depth in the tissue. The present results, estimating a depth of 0.5 mm for 480 nm, may encourage clinicians to use this wavelength for sterilizing infected wounds. It is therefore recommended to illuminate infected wounds with 480 nm for killing the pathogens prior to 730 nm light for stimulating skin cells growth and wound closure. The potential of 480 nm to enhance NO formation for increasing blood circulation in wounds should also be considered. In conclusion, the photon migration model substantiates theoretically the clinical routines of wound healing using lasers and light emitting diodes (LEDs) in the red-nir region [30,31] and suggests the 730 nm to be the optimal wavelength for enhancing wound closure. 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