1.

Introduction

Diffuse optical spectroscopic (DOS) techniques have been widely used to obtain in vivo tissue optical properties.^1,2 Using light transport models in the temporal and spatial domains, these techniques can quantify tissue absorption and scattering.^3–5 Specifically, DOS techniques quantify the wavelength-dependent tissue reduced scattering (${\mu }_s^{\prime }$) and absorption (${\mu }_a$) coefficients, which can be used to deduce subsurface structural and functional information. Recently, researchers have used ${\mu }_s^{\prime }$ to noninvasively assess wound healing.^6–8 In a porcine burn model, we previously showed that ${\mu }_s^{\prime }$ may accurately predict burn severity and skin wound healing capabilities.^9–11 Although these results showed promise for a potential new approach to rapidly assess burn severity and prognosticate wound healing, translating this technique to human subjects necessitates an understanding of baseline ${\mu }_s^{\prime }$ values in human skin. Thus, it is important to document ${\mu }_s^{\prime }$ values of normal skin at commonly used DOS wavelengths (visible- and near-infrared) for various anatomical locations and levels of pigmentation.

Prior to DOS, many studies have contributed to documenting human skin optical properties through in vitro and ex vivo measurements using integrating spheres. In 2011, Bashkatov et al.¹² thoroughly catalogued many of these contributions in their review work of in vitro, ex vivo, and in vivo optical properties of human skin, adipose, and muscle. These studies offered valuable insights toward complete characterization of human skin optical properties, for both whole skin and separated epidermis, dermis, and adipose layers. However, the reported values from these ex vivo measurements are not necessarily representative of in vivo tissues.^13,14 Previous in vivo studies to quantify ${\mu }_s^{\prime }$ of healthy skin in different anatomical locations have employed several different DOS techniques. Doornbos et al.¹⁵ utilized a fiber-based spatially resolved diffuse reflectance spectroscopy system to obtain in vivo optical properties of human skin and the underlying tissue. Tseng et al.^16,17 applied steady-state frequency domain photon migration to perform highly localized measurements of ${\mu }_a$ and ${\mu }_s^{\prime }$ of in vivo volar forearm, palm, dorsal forearm, and upper inner arm for human subjects across a range of Fitzpatrick skin types. In 2015, Saager et al.¹⁸ compared multiphoton microscopy and spatial frequency domain spectroscopy for measurement of melanin and reduced scattering on dorsal forearm and volar upper arm regions of 12 subjects of various skin types. In a study on volar forearm of 1765 Caucasian subjects (i.e., skin types I and II), Jonasson et al.¹⁹ obtained scattering parameters over a range from 475 to 850 nm using a commercial diffuse reflectance spectroscopic system. Kono et al. used reflection spatial profile measurement to measure optical properties at 450 to 800 nm and 950 to 1600 nm for 198 subjects on the inner forearm, cheek, and dorsal hand between thumb and forefinger.²⁰ However, these studies had two limitations: (1) they only covered a small range of anatomical locations for measurements of scattering properties and (2) the measurement systems were restricted to point-based or single-line measurements that required multiple measurements to characterize the heterogeneity of large regions on the body. In summary, clinical translation of DOS requires a broader characterization of in vivo human skin that spans multiple anatomical locations and pigmentation levels, while also accounting for the heterogeneous nature of each sampled region.

In this study, we employ spatial frequency domain imaging (SFDI) to characterize and document ${\mu }_s^{\prime }$ of normal skin for 15 subjects with various pigmentation levels (Fitzpatrick types I to VI²¹) at 10 anatomical locations. SFDI is a noncontact, wide-field DOS imaging technique that uses spatially modulated illumination in combination with models of light–tissue interaction to determine optical properties of in vivo tissue.^5,22,23 Dognitz and Wagnieres²² first developed and used a variation of SFDI to obtain in vivo skin optical properties at 400, 500, and 700 nm. Cuccia et al.^5,23,24 further developed the technique to expand the imaging spectrum to include near infrared wavelengths and enabled clinical translation of SFDI to skin ulcer imaging. Here, we document ${\mu }_s^{\prime }$ values across all subjects and anatomical locations at imaging wavelengths, ranging from visible to near-infrared. These measurements were derived from the semi-infinite homogeneous model described previously.⁵ We then compare ${\mu }_s^{\prime }$ values between subjects at each wavelength and identify 851 nm as the wavelength with the least variation in ${\mu }_s^{\prime }$ between subjects. We posit that this result is due to melanin being highly absorbing at visible wavelengths and localized in a thin layer at the base of the epidermis, which leads to a confounding effect in separating ${\mu }_a$ and ${\mu }_s^{\prime }$ in the visible spectrum for subjects with darker skin. The decreasing intersubject ${\mu }_s^{\prime }$ variation with increasing wavelength suggests that pigmentation effects on ${\mu }_s^{\prime }$ determined by SFDI are the least at longer wavelengths (i.e., near-infrared and beyond). Finally, we show that baseline ${\mu }_s^{\prime }$ values vary with anatomical location, using 851 nm (where absorption from melanin is the lowest) as the wavelength for this analysis. This study serves as the first report for categorization of normal human skin scattering properties across multiple skin types and anatomical locations using SFDI. These findings are important for establishing the natural variation in baseline ${\mu }_s^{\prime }$ that must be accounted for when DOS techniques are translated to a clinical setting for applications such as burn and wound healing triage.

2.

Materials and Methods

2.2.

Subjects

Subjects ($N=15$; 8 male and 7 female) were recruited and imaged under Institutional Review Board (IRB) protocol (IRB# 2011-8370). Subjects had skin types ranging from I to VI on the Fitzpatrick scale and no known dermatological complications. The majority of the subjects were young adults. Twelve subjects were in the age range of 18 to 35 years, whereas three were in the range of 36 to 55 years. During recruitment, we took care to ensure that subjects were distributed as evenly as possible across a wide range of skin types. However, we did not perform any a priori analysis to predefine the exact number of patients of each skin type to recruit. Measurements were obtained at 10 anatomical locations on each subject. Locations included cheek, ventral forearm, dorsal forearm, shin, palm, lower back, and chest (near collar bone), which are common areas for burn injuries. Measurements were also taken of the forehead, upper arm (bicep), and posterior neck (near the hairline). For regions not located on the midline, such as cheek, arm, and shin, we chose to image the subject’s dominant side. Fitzpatrick skin types were determined using subject surveys (Table S4 in the Supplemental Materials) and clinically verified by Dr. Sharif. In order to supplement the low-quality webcam images that are captured by the commercial SFDI device, color images were taken prior to each measurement, using a digital camera (NEX-3, Sony Corporation of America, New York, New York). Instrumentation and measured anatomical locations are shown in Fig. 1.

Fig. 1

(a) Cart-based SFDI instrument, OxImager RS™ (Modulim, Inc., Irvine, California), comprised of eight discrete LED light sources (471 to 851 nm) modulated at five spatial frequencies (0 to $0.2{\mathrm{mm}}^{-1}$). (b) Imaged anatomical locations.

3.

Results

3.1.

SFDI Measurements of Reduced Scattering Coefficients at Each Anatomical Location for Visible to Near-Infrared Wavelengths

To illustrate the variation in ${\mu }_s^{\prime }$ with wavelength and skin type, Figs. 2(a) and 2(b) show the ${\mu }_s^{\prime }$ spectra (averaged over an ROI) of the dorsal forearms and palms, respectively, for all 15 subjects, classified by Fitzpatrick skin type. Representative color images and ${\mu }_s^{\prime }$ maps of the dorsal forearm and palm of subjects with various Fitzpatrick skin types are shown in Figs. 2(c) and 2(d), respectively. Overall, at shorter wavelengths, subjects with Fitzpatrick skin types indicative of lower pigmentation (I, II, and III) had higher measured ${\mu }_s^{\prime }$ than subjects with skin types corresponding to more pigmentation (IV and V/VI), likely due to confounding effects from absorption of light by melanin.

Fig. 2

(a) ${\mu }_s^{\prime }$ distribution (471 to 851 nm) for the dorsal forearm across all wavelengths of all 15 subjects shown with their Fitzpatrick skin types. (b) ${\mu }_s^{\prime }$ distribution (471 to 851 nm) for the palm. (c) Representative color and ${\mu }_s^{\prime }$ image examples of the dorsal forearm on subjects of Fitzpatrick types I, III, and V/VI. Scale bar = 1 cm. (d) Representative color images and ${\mu }_s^{\prime }$ maps of the palm. Scale bar = 1 cm.

The intersubject coefficients of variation for ${\mu }_s^{\prime }$ across all 15 subjects were calculated for each anatomical location at each wavelength (Table S3 in the Supplemental Materials). These coefficients decreased with increasing wavelength for all 10 anatomical locations (0.554 to $0.682{\mathrm{mm}}^{-1}$ at 471 nm; 0.0789 to $0.111{\mathrm{mm}}^{-1}$ at 851 nm). These values showed the least intersubject variation at 851 nm. The decrease in intersubject coefficient of variation of ${\mu }_s^{\prime }$ with increasing wavelengths coincides with the monotonically decreasing eumelanin extinction coefficient.^26,27 This result suggests that variation in ${\mu }_s^{\prime }$ at shorter wavelengths is largely due to the inability of the semi-infinite homogeneous light transport model to adequately extract optical properties in subjects with darker skin types. In Table 1, we show the intersubject ${\mu }_s^{\prime }$ means and standard deviations at 851 nm, the measured wavelength that we believe is the least confounded by pigmentation.

It should be noted that the palm also showed the decreasing trend in coefficient of variation for ${\mu }_s^{\prime }$ with wavelength, but the decrease was less pronounced (0.168 to $0.0596{\mathrm{mm}}^{-1}$ over the same range of wavelengths; Table S3 in the Supplemental Materials). This is most likely due to the palm possessing the lowest melanin concentration in comparison to other anatomical locations.^28,29 Thus the palm ${\mu }_s^{\prime }$ values are least confounded by pigmentation. Intersubject means and standard deviations of ${\mu }_s^{\prime }$ and ${\mu }_a$ at each anatomical location are documented in Tables S2 and S4 in the Supplemental Materials. Figure S1 in the Supplemental Materials shows ${\mu }_a$ values for each patient measured from the dorsal forearm.

3.2.

Interanatomical Location Variations of ${\mu }_s^{\prime }$ Among all Subjects at 851 nm

Based on the low intersubject variation in ${\mu }_s^{\prime }$ reported above, we chose 851 nm as the wavelength of interest for documenting skin scattering properties at the 10 anatomical locations for all 15 subjects. Table 2 shows the results of the one-way ANOVA test of comparing all 10 locations. Table 3 shows $p$ values obtained from Tukey’s tests between ${\mu }_s^{\prime }$ values at 851 nm for each pair of anatomical locations. Representative ${\mu }_s^{\prime }$ maps with ROIs are shown in Fig. 3, and average ${\mu }_s^{\prime }$ values at 851 nm for all subjects are shown in Fig. 4.

Fig. 3

Examples of ${\mu }_s^{\prime }$ maps from each anatomic location at 851 nm for a subject of Fitzpatrick skin type I. The 1-cm2 ROIs used for analysis are shown in white squares.

Fig. 4

Box and whisker graph of ${\mu }_s^{\prime }$ values measured at 851 nm from all 15 subjects for each anatomical location.

4.

Discussion

In this study, we used SFDI to obtain reduced scattering coefficient values (${\mu }_s^{\prime }$) of normal skin at 10 anatomical locations of 15 subjects ranging from Fitzpatrick skin types I to VI (Table 1). Our data showed lower ${\mu }_s^{\prime }$ values for subjects with higher Fitzpatrick skin type, which agrees with existing literature. Specifically, both Saager et al.¹⁸ and Jonasson et al.¹⁹ noted a decreasing ${\mu }_s^{\prime }$ for subjects with higher melanin fraction compared to ${\mu }_s^{\prime }$ measured in subjects having low melanin fraction. Tseng et al.^16,17 attributed this decrease in ${\mu }_s^{\prime }$ to limited probing depth due to increase in absorption by melanin, leading to fewer photons reaching the collagen and elastin matrix in the dermis, which can contribute strongly to ${\mu }_s^{\prime }$ values in the near-infrared wavelength range.

However, it must be noted that this substantial decrease in ${\mu }_s^{\prime }$ value with increasing pigmentation at visible wavelengths can also be attributed to the confounding effects of melanin’s absorption in the homogeneous processing model. We calculated depth penetration using the diffusion equation with a homogeneous tissue model where the input ${\mu }_a$ and ${\mu }_s^{\prime }$ values were obtained from SFDI.⁵ Measurements at visible wavelengths are interrogating small volumes and would be more affected by variation in melanin concentration existing in the epidermis. For longer wavelengths (i.e., near-infrared region), the penetration depth often surpasses the typical epidermal thickness of human tissues, which ranged from 100 to $150\mu \mathrm{m}$.³⁰ Examining the penetration depth for the planar ($0.00{\mathrm{mm}}^{-1}$) frequency of the dorsal forearm at 851 nm shows deeper mean penetration for subjects with lower Fitzpatrick skin rating (Fig. 5). However, there was no decrease in ${\mu }_s^{\prime }$ values for subjects having darker skin at 851 nm for the dorsal forearm [Fig. 2(a)]. This can be attributed to longer wavelengths probing further into the tissue due to: (1) the reduction in extinction coefficient of melanin along with (2) a decrease in ${\mu }_s^{\prime }$ at longer wavelengths. Such increase in probing depth substantially extends the interrogating volume past the localized melanin layer, minimizing its confounding effects on the tissue’s overall optical properties.

Fig. 5

(a) Estimated penetration depth calculated for the planar ($0.00{\mathrm{mm}}^{-1}$) spatial frequency from dorsal forearm measurements, at 471 nm for subjects with Fitzpatrick scores I and II and (b) at 851 nm for all subjects. For 471 nm, only subjects with Fitzpatrick scores of I and II were chosen due to their optical properties being least confounded by existing melanin concentrations.

We observe a convergence of ${\mu }_s^{\prime }$ values for all skin types at 851 nm [Figs. 2(a) and 2(b)]. This result is also seen in the decrease in intersubject coefficients of variation for ${\mu }_s^{\prime }$ values with increasing wavelength [Fig. 2(a)]. Saager et al.¹⁸ and Jonasson et al.¹⁹ also found this convergence of scattering properties for different skin types at longer near-infrared wavelengths in their respective studies. Since melanin absorption decreases with wavelength in this regime [Fig. 2(c)], longer wavelengths tend to minimize inaccuracies during the fitting process for ${\mu }_a$ and ${\mu }_s^{\prime }$ when using a semi-infinite homogeneous model. The low coefficient of variation across skin types for the palm [Fig. 2(b)] further suggests a minimal effect from the contribution of melanin toward ${\mu }_s^{\prime }$ variations. Previous studies have shown that the palm’s fibroblasts secret DKK1, an inhibitor of $\mathrm{Wnt}/\beta$-catenin pathway preventing growth and functionality of melanocytes.^28,29

We have also shown that SFDI measurements of ${\mu }_s^{\prime }$ values varied among different anatomical locations, even at 851 nm, for all subjects (Table 3, Figs. 3 and 4). The difference in scattering properties among anatomical locations has been previously attributed to anatomical structural variations,^16–18,20 including skin thickness, collagen structures, and mitochondrial density.³¹ For all 15 subjects, we observed highest ${\mu }_s^{\prime }$ values for cheek and forehead (averaged 1.52 and $1.64{\mathrm{mm}}^{-1}$, respectively, in Fig. 3). This agrees with the findings of Takema et al.³² Specifically they found that facial skin regions, because they are constantly exposed to sunlight, increase in thickness over time in comparison to skin on the ventral forearm. The ${\mu }_s^{\prime }$ values that we measured on the ventral forearm at 851 nm were $1.45\pm 0.115{\mathrm{mm}}^{-1}$, in comparison to $1.13\pm 0.27{\mathrm{mm}}^{-1}$ at 850 nm reported for a Swedish cohort of 1734 subjects.¹⁹ The discrepancy between results may be a consequence of the different methods of calibration used by the different groups. Although we use a physical tissue-simulating phantom (with optical properties measured using an integrating sphere) as a calibration to obtain the diffuse reflectance, Jonasson et al. normalized an Inverse Monte Carlo modeled spectrum with the average measured spectral intensity to avoid the need for absolute calibration of the intensity recorded by their spectrometers. Furthermore, we also imaged skin for all Fitzpatrick skin types. This will contribute to greater variation in our data relative to the Swedish cohort.¹⁹

Finally, it should be noted that the ${\mu }_s^{\prime }$ spectra for Fitzpatrick skin types I and II appear typical for biological tissue and can be described by Rayleigh and Mie scattering.^19,31,33,34 However, our measured ${\mu }_s^{\prime }$ from darker skin (Fitzpatrick skin types III to VI) tended to increase with wavelength for almost all anatomical locations except the palm. We attribute this observation to the high absorption of visible light by melanin found only in the epidermis. The localization of melanin to the epidermis contributes to a highly inhomogeneous depth-resolved absorption profile. Our simple semi-infinite homogeneous model cannot resolve such complex geometry.³⁵ Evidence of this limitation can be seen from an examination of the data obtained for the palm. The palm has a very low melanin fraction for all skin types. For palm skin, we observed a decrease in ${\mu }_s^{\prime }$ spectra with increasing wavelength [Fig. 2(b)]. We are currently investigating the effects of melanin confined to the epidermis on SFDI derived optical properties determined from a semi-infinite homogeneous model and developing new models to account for such effects.

5.

Conclusion

In this study, we have for the first time documented the reduced scattering coefficient properties ${\mu }_s^{\prime }$ of normal skin at 10 anatomical locations, for subjects having pigmentation variations across all Fitzpatrick skin types (i.e., I to VI), using SFDI. Examining the measured ${\mu }_s^{\prime }$ at an anatomical location (i.e., dorsal forearm) at visible wavelengths showed a decreasing trend with higher Fitzpatrick skin type. However, there were no significant differences of this kind observed between any of the skin types seen at the longest wavelength measured (851 nm). Furthermore, significant differences in measured ${\mu }_s^{\prime }$ at 851 nm were observed between different anatomical locations. These findings regarding the reduced scattering properties across various anatomical locations of subjects of all Fitzpatrick skin types will aid in establishing baseline SFDI measurement for future clinical studies.