3.1.

Variability Among Absorption and Scattering Coefficients: Individual Publications

The graphs in Fig. 4 show the variability among the published absorption and reduced scattering coefficients for the different FST groups compiled from 12 different publications. As well as the range in spectra used to determine the optical properties of skin, these data also show that there is variability in sample number among the published in vivo data and that data are limited for FST V–VI. In total, absorption measurements were taken from in excess of 58 subjects for FST I–II (two papers do not state the sample number) with a modal sample number of 3.^37–43 For FST III–IV, there were a total 301 subjects with one paper contributing data measured on 198 recruits and another measured 71 recruits.^{39,41,42,44–47} The modal subject number was 6. Only three papers detailed the optical properties of FST V–VI from a total of 12 subjects and the average sample number was 4.^39,41,42 Scattering measurements for FST I–II were from in excess of 1749 subjects, with one publication contributing data from 1734 subjects.^37–41,48 Published scattering data were available from 289 subjects for FST III–IV.^{39,41,44–47} Only two manuscripts described scattering data for FST V–VI from a total of seven recruits.^39,41

Fig. 4

Variability in absorption coefficients (a)–(c) and reduced scattering coefficients (d)–(f) among the published in vivo data with skin type. The number of subjects measured for each publication is shown as “$n$.” *Note that the data from Kono and Yamada⁴⁴ were converted to reduced scattering using an average $g$ taken from the publications of 0.84. **Data from forearm unavailable. ***Note that data measured by Saager et al.³⁹ for FST III and FST IV have been averaged to provide an FST III–IV measurement. Due to the reduced scattering scale distortion introduced by outliers, we point out that the distribution of reduced scattering data from Saager et al.³⁹ ranges from 2.27 to $1.18{\mathrm{mm}}^{-1}$, 1.94 to $1.042{\mathrm{mm}}^{-1}$, and 1.56 to $1{\mathrm{mm}}^{-1}$ for FST I–II, III–IV, and V–VI, respectively.

Variability between the published data within each FST group at given wavelength is huge with the average difference between the maximum and minimum reduced scattering coefficients across the measured wavelengths being 98%, 25%, and 31% for FST I–II, III–V, and V–VI, respectively. Average differences between the maximum and minimum reported absorption coefficients were 81%, 51%, and 38% for FST I–II, III–V, and V–VI, respectively.

Each author took measurements in different spectral bands, making reliable comparisons between the data difficult. The differences and trends among the publications are discussed here in further detail.

Tseng et al.⁴¹ split subjects into three skin type groups of six subjects based on melanin content using diffuse reflectance spectroscopy took measurements in the wavelength range 500 to 1000 nm. A decrease in absorption as the wavelength increased from 500 to 600 nm of twofold was observed for both skin types I–II and III–IV. The optical properties of FST V–VI were not measured in this region because this skin type was too absorbing for their measurement method. The mean absorption coefficients for FST III–IV were greater between 500 and 600 nm (the region coinciding with high melanin absorption) than for FST I–II. In the 600 to 800 nm spectrum, scattering by FST V–VI was greater than the other FSTs. Tseng et al.⁴¹ suggested that, since melanosomes (up to $1.44\mu \mathrm{m}$ in diameter^34,35) scatter light and scattering power is related to dimensions of the scatterer, the density of these in FST V–VI may contribute to increased scattering measurements observed in this study among this group.⁴¹

Measurements by Tseng et al.⁴¹ show that in the near-infrared (NIR) region of the spectrum, FSTs I–II and III–IV had very similar absorption and scattering properties. For FST V–VI absorption and scattering converged with the other skin types at 850 and 750 nm, respectively, indicating that the NIR is likely to be the region where FST may have less impact on optical properties. The differences between skin types I–II and V–VI were shown to be significant for both absorption and scattering in 600 to 850 nm and 600 to 750 nm ranges, respectively. A small peak in absorption was observed at $\sim 970\mathrm{nm}$ for all skin types corresponding with water absorption.

Bosschaart et al.⁴² took measurements in the visible spectrum (400 to 600 nm) of skin from neonates aged between 6 and 28 days with varying skin pigmentation at four different locations. In neonates, skin location did not have a significant effect on optical coefficients, but absorption was shown to increase significantly in skin type V–VI compared with skin type I–II and III–IV patients with time after birth. The data shown in Fig. 4 are taken from the median values for all measurements taken independent of time after birth (as shown in figure 7 from Ref. 42). Coefficients ranged among the grouped skin types between 0.02 to $1.25{\mathrm{mm}}^{-1}$ for absorption and 1 to $2.8{\mathrm{mm}}^{-1}$ for scattering.

Although the data gathered by both Bosschaart et al.⁴² and Tseng et al.⁴¹ only overlapped in the 500 to 600 nm spectrum, the variability in resulting data is exemplified by coefficient differences for each skin type in this region. At 600 nm, absorption data published by Bosschaart et al.⁴² are progressively greater with increasing skin pigmentation than Tseng et al.’s⁴¹ data, being 25%, 45%, and 77% greater for FST I–II, III–IV, and V–VI, respectively. At 500 nm, the differences are more marked, with Bosschaart et al.’s⁴² data being 73% and 77% greater than Tseng et al.’s⁴¹ for FST I–II and III–IV, respectively.

To quantify melanin in the epidermis of different FSTs, Saager et al.³⁹ measured the optical properties of the dorsal forearm for 12 subjects ranging in skin type and including FSTs I, II, III, IV, V, and VI where the number of subjects for each was 1, 3, 5, 1, 1, and 1, respectively. Although their data followed the same general trends as that of other authors, the optical properties were only measured up to 900 nm at which point absorbance by all skin types was similar. However, the optical properties of FSTs I and V were not described by these authors. Between 400 and 900 nm, absorbance for FST VI was greater than the remaining skin types. FSTs II and III had the least absorption, with both spectra being similar and FST IV absorption was between that of the palest and darkest skin. Scattering for FST II was greatest and least for skin type VI up to 800 nm; beyond this all skin types had similar spectra. Absorption by the darkest skin was similar to that measured by Bosschaart et al.,⁴² in the 450 to 600 nm spectrum and similar to measurements made by Tseng et al.,⁴¹ beyond 850 nm. However, unlike Tseng et al.,⁴¹ Saager et al.³⁹ measured a sharp decrease in absorption between 600 and 850 nm. Scattering measurements by Tseng et al.⁴¹ for FST V–VI were greater than those measured by Saager et al.³⁹ by twofold at 600 nm down to 18% at 900 nm.

The reduced scattering and absorption coefficients were determined between 471 and 851 nm by Phan et al.⁴⁹ for 15 subjects of all FSTs at varying locations. These data were not plotted in Fig. 4 because the peaks usually associated with the skin chromophores were not observed; however, the measurements were taken from one subject from FST I, three from FST II, six from FST III, three from FST IV, and two from FST V–VI. Absorption tended to decrease with increasing wavelength in all skin types. In general, subjects with skin types I and II had the least absorption in the visible spectrum. Subjects with skin types IV, V, and VI had the greatest absorption in the visible spectrum although there was some overlap with the data for skin type III.

The published data plotted in Fig. 4 suggest that absorption is generally greater for FST III–IV than I–II in the 400 to 450 nm spectrum, and in this region absorption by FST V–VI is less than for either of the lighter skin type groups (i.e. FST I-II and III-IV). It seems unlikely that absorption by FST V–VI would be less than by lighter FSTs due to the expected absorption by melanin and therefore differences may be attributed to author methods, sample numbers and locations, and potentially to determination of FST being subjective. Among the published data, only three authors produced data associated with FST V–VI compared with 15 datasets associated with the paler FSTs; this may be due to the prevalence of skin cancers in individuals of this skin type making it a research focus, or the country in which the research is taking place. Some studies with large sample numbers have been associated with the paler FSTs, for instance, the maximum sample number for FST V–VI is 6.^{41,42,44,46,48} For these reasons, we question the applicability of these data to skin types containing more melanin in their epidermis, i.e., FST V plus. Finally, not all useful optical properties have been measured at a wide range of wavelengths for all skin types in similar locations and there is variability in the wavelength ranges used by different authors and the measurements taken.

3.2.

Variability Among Absorption and Scattering Coefficients: Published Values Combined

Figure 5 shows a comparison between absorption and reduced scattering coefficients for averaged total published data from FST I–II, III–IV, and V–VI. Data from one publication by Svaasand et al.⁴⁰ were not included in the average measurements for the reasons discussed by Lister et al.⁵⁰ Briefly, reduced scattering and absorbance measurements were only taken for FST I–II and are considerably higher than those measured by other authors being $\sim 80\%$ greater than the average absorption measured by other authors and $\sim 97\%$ greater than the average reduced scattering across the measurement spectrum, as can be seen in Fig. 4. Therefore, although the data included are very variable, this dataset is an extreme outlier and would potentially unrealistically skew the average data and hide differences that might be possible to determine between FSTs. FST V–VI absorption coefficients are $\sim 7\%$ to 74% greater than those for FST I–II in the 400 to 1000 nm spectrum, the difference being least at 440 nm (6.5%) and greatest at 640 nm (74.4%). In the NIR absorption, coefficients start to converge but tend to increase in value with increasing wavelength. However, data are limited in this region, and beyond 1000 nm measurements were only available for FST III–IV. Therefore, it would be useful to gather further data for wavelengths beyond 1000 nm to determine whether the optical coefficients for the varying FSTs continue to converge with increasing wavelength. Beyond 1000 nm, data published for FST III–IV show absorption peaks associated with water.

Fig. 5

Differences between averaged published (a) absorption and (b) reduced scattering coefficients for FST I–II,^37–43,48 III–IV,^{39,41,42,44–47} and V–VI.^39,41,42

Reduced scattering coefficients decrease with increasing wavelength for all FSTs. The data suggest that there are fewer scattering events in more melanized skin types at visible wavelengths (400 to 600 nm), where the reduced scattering coefficients for FST V–VI were up to twofold less than for FST I–II. The observed scattering differences in the visible spectrum are most likely to be due to absorption differences among skin types in this spectrum affecting scattered photon number and detection, i.e., FST V–VI absorbs more light in the visible region, therefore, proportionally fewer photons are available to be scattered and detected than in FST I–II. However, it is difficult to draw any conclusions from this due to the large variability in the published data contributing to this. It should also be noted that the data contributing to this region of the graph come from a single publication and sampled from only one subject leading to a marked skew in the data. Beyond 600 nm, there was a tendency toward convergence of the reduced scattering coefficients for FSTs III–IV and V–VI and between 600 to 1000 nm the reduced scattering coefficients associated with FST I–II were 22% to 48% less than for the darker FSTs. Reduced scattering coefficients were most similar between the palest and darkest FSTs at 600 nm, where the difference was 20%.

3.3.

Transport Mean Free Path of Light Through the Skin

To put the values presented in Sec. 3.2 into context, we consider the maximum depth at which it would be possible to focus light into skin for different types and hence how deep into skin an image can easily be formed (see Fig. 6). The transport mean free path (TMFP) was used to provide a measure of this depth, where $\mathrm{TMFP}=1/{\mu }_s^{\prime }$ and ${\mu }_s^{\prime }$ is the reduced scattering coefficient given by ${\mu }_s^{\prime }={\mu }_s(1-g)$, where ${\mu }_s$ is the scattering coefficient of the material and $g$ is the anisotropy factor representing the average cosine of the scattering angle.⁵¹ The TMFP predicts how far light will travel through a material before it becomes diffuse with the direction of propagation no longer resembling the starting direction of the beam. At this depth, it is virtually impossible to focus the light with standard optics and without the use of wave-front correction or adaptive optics.

Fig. 6

Average TMFP (mm) associated with each FST group.

The results gathered from the published data and analyzed here suggest that beyond 600 nm the depth at which an image could potentially be formed increases with increasing wavelength and is greatest for FST I–II at 980 to 1000 nm being 1.37 mm, for FST III–IV at 1470 nm being 1.27 mm, and for FST V–VI at 940 nm being 0.83 mm. For comparison, the maximum TMFP for FST III–IV in the 400 to 1000 nm range is 0.82 mm at 950 nm. Because the TMFP is related to scattering, and the NIR is the region where scattering is least (Sec. 3.2), this is the region where light is likely to travel the furthest distance into the skin. Like scattering, differences between the TMFP for FST III–IV and V–VI are least in the NIR spectrum. The observed TMFP magnitude in the 400 to 600 nm spectrum is unlikely to be realistic as described in Sec. 3.2. There is the biggest difference between FST V–VI and FST I–II between 960 to 1000 nm where the TMFP is $\sim 90\%$ greater for FST I–II than the darker FSTs. Accounting for the above, in the 400 to 1000 nm spectrum, it is likely that light of between 940 and 1000 nm may provide a useful wavelength for imaging skin of different FSTs at depth. However, there is a small peak in absorption at these wavelengths as shown in Fig. 5, and because TMFP calculations do not take absorption into consideration it is likely that the amount of light reaching these depths will be affected. Absorption coefficients are, however, at least an order of magnitude less than reduced scattering coefficients at these wavelengths so absorption may have little effect on the distance light can travel through the skin or the amount of light that can reach the depths. This is discussed further in Sec. 3.4

We are unable to determine whether the TMFP may increase further beyond 1000 nm as shown for FST III–IV because comparative data for FST I–II and V–VI do not exist at these wavelengths among the published data used for this review. It would, therefore, be useful to gather data beyond a wavelength of 1000 nm for all FSTs to determine if the optical coefficients decrease with increasing wavelength and to therefore determine if the maximum imaging depth can be extended for longer wavelengths.

The average TMFP across each FST of photons travelling through the skin ranges from 0.68 to 0.81 mm across the measurement spectrum depending upon FST. Using the published data, the average TMFP is least in FST V–VI samples and greatest in FST III–IV in the 400 to 1000 nm spectrum. It should be noted that TMFP does not consider absorption and a beam travelling through FST V–VI skin would likely be more affected by attenuation due to absorption than FST I–II. In practice, the power of the light source could be increased to account for absorption losses but a consequence of this would be an increase in the risk of photodamage to the tissue.

3.4.

Transmission of Light Through the Skin

Estimating how much light will reach a certain depth into the skin is important for medical treatments that use light and photons such as PDT. With scattering present it is accepted that beyond the mean free path a form of the Beer Lambert law can be written with the attenuation written as $e^{-{\mu }_s^{\prime }l}$, where depth is $l$ and the reduced scattering coefficient is ${\mu }_s^{\prime }$.^52,53 When absorption and multiple scattering is present, the situation is more complicated because the total propagation path of the detected photons is generally larger than $l$ (the penetration depth). Monte Carlo simulation can be used to simulate the path length of the photons arriving at a particular detection position, which can be recorded. In practice, the additional propagation length is a function of the scattering coefficient, the penetration depth, and the detection mechanism. This is discussed by Sassaroli and Fantini,⁵⁴ and elsewhere.

For our purposes, to come up with some plausible values, we consider the case where we wish to focus the optical beam to a well-defined position as is the case with, for example, PDT or nonlinear microscopies (such as two photon harmonic, etc.). In this case, it is reasonable to say that the light available at a certain depth to form a focus has suffered far less scattering than a “typical” photon. Bearing this in mind, the intensity of light reaching a depth, $l$, into a medium with a reduced scattering coefficient of ${\mu }_s^{\prime }$ and an absorption coefficient of ${\mu }_a$ can be expressed as $I_{\text{out}}=I_{\text{in}}{e}^{-({\mu }_a+{\mu }_s^{\prime })l}$, where $I_{\text{in}}$ is the input intensity and $I_{\text{out}}$ is the intensity at a depth of $l$. In practice, this is represented as the percentage of light that has been transmitted or has reached $l$. We acknowledge that this is a somewhat ad hoc expression and will tend to overestimate the amount of light reaching a particular position, it will, however, give a useful ballpark and a method for considering the challenges of certain technologies with skin color.

The percentage of light propagating through given depths using averaged published data for each FST was calculated. Figure 7(a) shows that at all wavelengths measured in the published data, more than 69% of light was transmitted beyond a penetration depth ($l$) of 0.1 mm, which corresponds to the average thickness of the epidermis.⁴³ However, the percentage of light transmitted beyond this depth was variable both for wavelength and skin type. In general, between 600 and 1000 nm light travels further into FST I–II skin than FST III–IV and V–VI skin according to TMFP calculations as shown in Fig. 6. Similarly, the percentage of light transmission was shown to be greatest for FST I–II in this spectral region. The maximum amount of light transmitted beyond 0.1 mm is 92.6% at 1000 nm, 88.7% at 1290 nm, and 88.2% at 940 nm for FST I–II, III–IV, and V–VI, respectively. The measured spectral range for FST III–IV skin is much broader than for the other two skin type groups; if the same range is considered (400 to 1000 nm) then the maximum light transmitted through this group is 88.2% at 950 nm. The percentage transmission for FST III–IV and FST V–VI is very similar to each other in the 600 to 940 nm wavelength range. The data contributing to the transmission calculations are extremely varied as shown in Fig 4; averaging of these is likely to have contributed to the similarity of the reduced scattering coefficients and therefore the percentage transmission, which is largely affected by transmission, in the 600 to 940 nm range.

Fig. 7

(a) Transmission of light at 0.1 mm* dependent upon skin type. (b) Transmission of light beyond 1.93 mm* dependent upon skin type. FST I–II, red; FST III–IV, green; FST V–VI, blue. *0.1 and 1.83 mm correspond to the average epidermal and dermal thickness respectively giving a total depth of 1.93 mm.⁴³

It might be expected that light transmission through all FSTs would be maximal at 1000 nm (and potentially even longer wavelengths if the data were available). However, due to the variability and limited nature of the data that make up the averaged coefficients used for this work, it is difficult to reliably and precisely extract information. Calculations suggest that 940 to 1000 nm may be a useful wavelength range for enabling most light to be transmitted through the skin as also suggested in Sec. 3.3. Despite the increasing absorption observed beyond 940 nm (described in Sec. 3.2), because scattering is least for all FSTs at these wavelengths, transmission is greatest in this spectrum. Although this region coincides with water absorption peaks, because absorption has much less of an effect than scattering, absorption by water may not have a significant negative effect on transmission.

Based on the published data, largest losses of light occur with increasing skin depth; across the 400 to 1000 nm spectrum on average $\sim 14\%$ of light is lost at a depth of 0.1 mm in FST I–II and by a depth of 1.93 mm 91% of the remaining light has been lost. In FST III–IV and V–VI skin, these losses are on average 18% and 16%, respectively, at 0.1 mm and 97% and 96%, respectively, by a depth of 1.93 mm. For FST III–IV beyond 1000 nm less light is lost than in the 400 to 1000 nm spectrum, with 12% average losses at a depth of 0.1 mm and 92% by a depth of 1.93 mm. Because data are not available for FST I–II and V–VI beyond 1000 nm, we are unable to determine whether less light is lost with increasing wavelength as shown for FST III–IV. To put these losses into context, if a biomedical technology requires 10 mW of light to be transmitted beyond a depth corresponding to the average combined depth of the epidermis and dermis (1.93 mm), the skin must be illuminated with $\sim 406.5\mathrm{mW}$ of light (this is based on FST III–IV, where average light losses are greater than FST I–II and V–VI in the 400 to 1000 nm spectrum). However, losses are wavelength dependent as well as FST dependent and since most light is lost at shorter wavelengths the power used to illuminate the skin would need to be much greater at these wavelengths (up to 18 W). By comparison, beyond 940 nm the power required to illuminate the skin and have 10 mW remaining beyond the epidermis and dermis is as low as 47 mW for FST I–II. However, despite being least at these wavelengths, more than twice this power would be required for FSTs III–IV and V–VI.

More than 1% of light was transmitted beyond the full depth of the skin (1.93 mm) at all wavelengths measured for FST V–VI. However, in the 600 to 900 nm spectrum, the percentage of light transmitted beyond this depth was greater for FST I–II than for FST V–VI. This coincides with the region discussed in Sec. 3.1 where elevated scattering by FST V–VI in the 600 to 800 nm region is attributed to the scattering power of melanosomes.⁴¹ For more than 1% of light to be transmitted beyond the skin, at least 78% of light must be transmitted beyond 0.1 mm [Fig. 7(b)]. For FST I–II and III–IV, this is most likely to happen beyond 600 nm. More than 78% of light was transmitted beyond 0.1 mm for FST V–VI at all wavelengths measured [Fig. 7(a)] and therefore, more than 1% of light has the potential to be transmitted beyond skin with a total depth of 1.93 mm, the average combined depth of epidermis and dermis described by Meglinski and Matcher.⁴³

Light transmission beyond 0.1 mm was the least at the shortest wavelengths and the minimal transmission reported for each FST in the visible wavelength range being 78.9% for FST V–VI (at 450 nm) compared with 71.2% for FST I–II (at 420 nm) and 69.2% for FST III–IV (at 420 nm). In the 400 to 1000 nm spectrum, transmission was greatest through all FSTs at wavelengths longer than 940 nm. At 940 nm, light is still travelling predominantly in the forward direction (as shown by TMFP calculations) at a depth of 0.1 mm and is at $\sim 88\%$ on average of its original intensity for all FSTs. At this wavelength and a skin depth of 1.93 mm, light transmission has reduced to between 18% and 23% in FST I–II compared with between 8% and 9%, and between 6% and 9% in FST III–IV and V–VI, respectively.

Nishimura et al.⁵⁵ described how the penetration depth of optical measurements changes significantly at wavelengths corresponding with strong water absorption at longer wavelengths. They found that the penetration depth is less than 1 mm in the 1400 nm water absorption band, which means that at this wavelength light would not penetrate beyond the skin.

By modeling photon transmission and using data taken from published results for the optical properties of the skin, Finlayson et al.⁵⁶ showed that 1% of light reached depths of 1.6 and 5 mm at wavelengths of 450 and 650 nm, respectively. Similarly, using Monte Carlo (MC) modeling, Ash et al.⁵⁷ showed that at wavelengths of 450 and 650 nm 1% of light has the potential to reach 1.6 and 4.75 mm, respectively.

Sordillo et al.⁵⁸ described the potential of longer wavelength light to follow a linear trajectory for a greater distance than shorter wavelength light, due to a reduction in scattering and absorption coefficients. For published data assembled by Shaw et al.,⁵⁹ it was shown that light in the 1000 to 2000 nm spectrum can increase penetration depth into biological tissues, improve image resolution, reduce tissue autofluorescence, and aid image-based diagnostics. For this reason, gaining a set of data relating to the optical properties of the skin beyond 1000 nm for different FSTs would be beneficial.

The results of analysis of the published data described in detail in this section are summarized in Table 2. Note that the wavelengths shown are those that are important for the optical medical technologies that are described in Sec. 4.

Table 2

Summary of the average absorption and reduced scattering coefficients (mm−1), calculated TMFP (mm) and percentage transmission at depths corresponding to the average epidermis (0.1 mm) and the combined epidermis and dermis (1.93 mm). Data are given at wavelengths important to OCT, PDT, and medical wearables.

4.1.

Optical Imaging Using OCT

Optical imaging is an important technique because it significantly reduces a patient’s exposure to ionizing radiation from, for example X-rays, and can therefore be used repeatedly for monitoring disease progression and treatment results.⁸¹ There are multiple forms of optical imaging and a recent review compared OCT, reflectance confocal microscopy, and multiphoton microscopy as techniques for noninvasive optical biopsy of in vivo skin of patients with pigmented lesions, such as melanoma and basal cell carcinoma (BCC).⁸² These all use wavelengths beyond 700 nm and therefore avoid maximal melanin absorption reducing the effect of skin color. Of these, OCT provides the greatest imaging depth and field of view.

OCT produces three-dimensional images noninvasively using low coherence interferometry to image within optically scattering samples, such as biological tissues. In OCT, the beam is split in two to give a reference beam, normally reflected off a mirror, and a sample beam, backscattered from the tissue of interest, and the interference between the two beams is measured. The use of interferometry means that photons that have been scattered multiple times and are incoherent are rejected, and so will not obscure the image by contributing to background noise. OCT is sensitive to scattering from tissues because coherence is essential for detection; OCT detects back-reflected light that is coherence matched with the incident light.^83,84 OCT imaging at depth is negatively affected by absorption with light absorption and attenuation in upper tissue layers leading to less light reaching lower layers. Light reaching these lower layers can also become attenuated on its return prior to detection.⁸⁵ Widefield imaging methods can only image samples to depths of $\sim 20\mu \mathrm{m}$; however, OCT can be used to image to depths of up to 2 mm with resolution smaller than $10\mu \mathrm{m}$.^83,86 OCT is currently unable to image much deeper than a few millimeters because beyond this depth the proportion of light that has not under gone multiple scattering events is too small to be detected.

OCT has been adopted most commonly for obtaining noninvasive high-resolution images to monitor eye health (e.g., macular degeneration, retinopathy, and optic nerve neuropathies). However, other clinical applications include obtaining detailed images of coronary arteries, guiding treatment, and monitoring stents in cardiovascular diseases; OCT in combination with endoscopy to detect and guide treatment of pancreatic cancer; as well as becoming more commonly used in dermatology over recent years including diagnosis and monitoring of skin disorders.^86–95

Because OCT resolution and imaging depth are affected by absorption, the melanin content of the skin and its effect on absorption need to be considered. In the field of dermatology, OCT is most extensively used in the diagnosis of BCCs; however, it currently does not have the accuracy to diagnose malignant melanomas.⁹⁶ It has also been shown to have potential for monitoring therapy for skin disorders such as scleroderma, psoriasis, eczema, and wound healing.^96–98 The skin layers from the stratum corneum to the upper dermis, and skin appendages and blood vessels can be clearly imaged.⁹⁵ Although OCT has been shown to accurately image the development of actinic keratosis (AK), a precursor to squamous cell carcinoma (SCC), in patients of FST I–III, the majority of studies do not mention skin color and this is likely due to the perception that AK only affects people with light skin.^99–104 However, although AK is more common in people with light-colored skin, it can be found in patients with all FSTs.¹⁰⁵

The effect of FST on OCT imaging has been rarely investigated. A 2019 publication suggested that image quality may become slightly lower with increasing pigmentation, but concluded that the difference in imaging quality and depth was not significant among all FST.⁸ Here, OCT with a center wavelength of 1310 nm was used and a depth of 1.3 mm below the skin surface was imaged. The authors suggest that despite having quantitatively lower quality, where light intensity at this depth was significantly less in dark skin than light skin, imaging is qualitatively indiscriminate for skin color. However, although melanin absorption should be minimal in this region, as described in Sec. 2, melanin may still have absorbing properties even at this wavelength.³¹ Ekelem et al.⁸ suggested that further work is required to study the effect of skin type on OCT imaging, and that their small sample size [16 subjects in total; FST I–III ($n=8$) and FST V–VI ($n=8$)] and sun-exposure of the subjects may have influenced results.

The effect of skin color on OCT has rarely been discussed making it difficult to draw conclusions. This is presumably because OCT uses longer wavelengths, which are supposed to be unaffected by absorption by melanin. However, the majority of studies have also been in patients with FST I–II; although OCT tends to use wavelengths beyond 1300 nm for skin imaging, darker skin has rarely been studied and it has been suggested that OCT image quality may be reduced in patients with darker skin but further investigation is required.⁸

4.2.

Photodynamic Therapy

PDT involves the use of light sensitive medications to destroy abnormal cells and can be used to treat some diseases of the skin such as actinic keratoses and BCC, warts and acne as well as macular degeneration of the eye, and cancers of the esophagus, mouth, and lung.⁶² For PDT to be effective, a clear understanding of the propagation of light through human tissue is required, which is affected by local absorption and scattering coefficients.³ In general, the scattering coefficients of tissues decrease monotonically with increasing wavelength; however, absorption coefficients vary greatly depending upon the absorbing chromophores present. For example, hemoglobin in the blood and melanin in the skin are strong absorbers of visible light, whereas water has multiple absorption peaks at infrared wavelengths. PDT can be applied either superficially to noninvasively treat accessible regions such as the skin; invasively, where superficial applications would have limited light penetration, to treat bulk tissues with optical fibers being placed directly in the tissue (interstitial treatment); or within a cavity such as the bladder via an endoscope.³

PDT requires that a photosensitive drug is activated in the presence of light and oxygen to produce reactive oxygen intermediates that irreversibly damage cells.⁶¹ This allows selective treatment; however, selectivity is affected by uptake of the photosensitive drug by target cells and its metabolism to an active form as well as by penetration of the light source. This is, in turn, affected by the optical properties of the tissue and, for treatment of a skin-disorder, the optical properties are affected by FST. Inflammatory hyperpigmentation is a side effect more commonly associated with the treatment of patients with FST IV–V and caused by excess synthesis and deposition of melanin enhanced by inflammatory responses.²⁰ Potential reasons for the common association of inflammatory hyperpigmentation with dark skin are increased oxidative stress and reduced vasodilation, which are improved with vitamin D (the intrinsic production of which is reduced in darker skin), as well as increased circulating inflammatory markers, which could lead to impaired melanin production and altered skin cell activity.¹⁰⁶

A wide variety of wavelengths are used for PDT. Blue light (450 to 495 nm) is used to maximize absorbance and longer wavelengths of light (beyond 620 nm) have best tissue penetration.^107,108 However, photosensitive drugs are only able to generate singlet oxygen, the most destructive reactive oxygen species, in the presence of light from 600 to 800 nm.¹⁰⁹

Two photosensitizing drugs are commonly used for PDT and the treatment of skin conditions; aminolevulinic acid (ALA) and its methyl ester cream [methyl aminolevulinate (MAL)]. Both are absorbed into the skin becoming converted to protoporphyrin IX with a major absorption peak in the blue light spectrum (410 to 420 nm).^61,110 However, due to penetration, red light (635 nm) is usually used to activate protoporphyrin IX for skin cancer treatments.¹¹¹ Because MAL is more lipophilic than ALA it is absorbed more deeply into the skin and is therefore commonly used in conjunction with red light, which penetrates more deeply into the skin than blue light.¹¹² Treatment with MAL-PDT has most commonly been associated with treatment of nonmelanoma skin cancers and, more recently, acne.^112,113 A further photosensitizing drug, chlorin, comprising of chlorin e6, chlorin p6, and purpurin 5, has been approved for use in Russia and Republic of Korea with absorption at 662 nm and has been shown to be less painful to patients than MAL and ALA treatments.^66,67,114 Because light penetration is limited in the visible range, these drugs are used for the treatment of superficial skin diseases with therapeutic depths of $<2\mathrm{mm}$; AK, for example, is usually $<300\mu \mathrm{m}$ deep.^115,116

Patients with FST IV–VI are at a greater risk of epidermal injury and hyperpigmentation or hypopigmentation due to PDT than those with light skin due to laser energy absorption by melanin at the wavelengths used for treatment.²⁰ Accurate determination of the individual’s skin properties and the associated absorption and scattering coefficients is required to reduce adverse effects by personalizing treatment regimens including reducing the doses of light, the strength of the photosensitive drug, and the incubation time of the drug on the skin.^65,117

For the treatment of acne, the total dose of red light needs to be adjusted for FST as discussed by multiple authors for patients of Asian origin.^65,118–120 However, there has been very little research into patients with the darkest skin types.¹²¹ Where PDT has been used to treat nonmelanoma skin cancers, including BCC, SCC, Bowen’s disease, and AK, the majority of the published work has studied efficacy in people with FST I–II.^{112,122–125} This is probably because skin cancers are more prevalent in pale-skinned populations with the people with the darkest skin being diagnosed, with approximately one twentieth of the cases of skin cancer of those with the palest skin.¹²⁶ Nevertheless, there are incorrect perceptions that skin cancers are unique to the lighter-skinned population and morbidity in darker-skinned people is greater due to later diagnosis, with 70% 5-year survival rate in dark-skinned people compared with 92% in lighter-skinned people.¹²⁷ However, incidence and morbidity are difficult to determine accurately due to limited availability of data in people with dark skin.¹²⁸ Farberg et al.¹²⁹ suggested that to evaluate the efficacy of PDT accurately, clinical studies should represent a greater diversity of FSTs, age groups, and anatomic locations.

PDT has mostly been investigated in light skinned patients and this is probably due to its use in skin cancer treatment and the increased likelihood of skin cancer in these patients.¹²⁴

4.3.

Medical Wearables and PO

Medical wearable devices are worn directly on the body to measure vital signs including heart rate, blood pressure, and oxygen levels. In healthcare, the most widely used method for optical sensing is photoplethysmography (PPG), which monitors blood flow in real time.¹³⁰ Differences in the interaction of light with human tissues at differing wavelengths is exploited to provide information including oxygen saturation, blood pressure, and more recently other cardiovascular-related diseases.¹³¹ The PPG signal can either be measured in transmission, where the light emission is physically opposite to the detection apparatus, or reflection, where the light emission and detection apparatus are on the same side as each other. In transmission, the signal from high capillary regions such as earlobes and fingers is most commonly measured. Light of 680 or 810 nm is used due to its penetration depth and ability to differentiate between oxygenated and deoxygenated blood. The light intensity transmitted through the skin will depend upon the volume of blood at the measurement site. For reflectance measurements used, for example, for heart rate monitoring, green light is usually used as this provides high quality PPG signals and is less affected by temperature changes and accuracy.¹³²

Currently, the most common form of PPG is PO, which works in transmission and uses two light sources to measure oxygen saturation of the blood.¹⁰ At the wavelengths used for PO, our calculations suggest that the TMFP for FSTs I–II, III–IV, and V–VI are 0.92, 0.61, and 0.64 mm, respectively. Oxygen saturation is estimated as a ratio of ratios of the minimum (DC: non-pulsatile ‘direct current’ blood component) versus peak (AC: pulsating ‘alternating current’ arterial blood components) light emanating from oxygenated blood, which transmits red light (622 to 780 nm), and deoxygenated blood, which transmits infrared wavelengths (780 to 2400 nm) as follows

A clear understanding of the optical properties of the tissues being measured is required to be able to interpret the results appropriately. A recent flurry in publications identifies that skin pigmentation affects the accuracy of PO.^11,134–140 However, issues were identified in 2007 by Feiner et al.,¹⁴¹ who suggested that PO overestimates oxygen saturation in darker skin tones and up to 10% differences in saturation estimates were observed among different instruments. Using the ratio of ratios of pulsatile light to total transmitted light at both red and infrared wavelengths oxygen saturation of arterial blood is estimated. In theory, because the measurements should only be dependent upon arterial saturation, the estimated oxygen saturation should be independent of FST and other variables.¹⁴¹ However, where oxygen saturation is out of the normal range, this does not seem to be the case and absorption by melanin in darker skin appears to affect outcomes. Feiner et al.¹⁴¹ showed that at low oxygen saturation, the instruments overestimated the saturation levels in intermediate and dark-skinned subjects. To obtain a stable result from back scattered light being detected in PO, not only the blood’s absorption needs to be considered but also that of the skin. The attenuation of incident light used in PO to measure oxygenated blood (light in the red spectrum) by melanin implies that PO may be less reliable in dark skin.

The results of Feiner et al.¹⁴¹ and the reason that skin type needs to be considered have been exemplified by the recent COVID-19 pandemic.^134,142,143 Black and ethnic minority groups had poorest outcomes from COVID-19 infection and this was compounded by the overestimation of oxygen saturation.^11,134 The situation where actual oxygen saturation was $<88\%$ in patients but their measured PO calculations were $>92\%$ occurred three times more in Black patients than white meaning that many Black patients were at increased risk of undetected hypoxemia.^135,136 A recent study comparing arterial blood gas (invasive) with PO showed that Black patients were at much higher risk of having occult hypoxemia (an arterial oxygen blood saturation of $<88\%$) due to overestimation of blood saturation by PO thereby limiting access to supplemental oxygen for Black patients.¹³⁷

At the wavelengths used for PO, our calculations suggest that the TMFP for FSTs I–II, III–IV, and V–VI are 0.92, 0.61, and 0.64 mm, respectively. Because depth penetration of light is likely to be less in individuals with darker skin tones, this may also be a contributing factor to the bias associated with PO.

Other wearable healthcare products used to measure heart rate and arrhythmia, for example, have also been shown to be limited by racial bias.¹⁴⁴ Colvonen et al.¹⁴⁴ expressed concern that as optical wearable devices are becoming more common in healthcare, and as evidence mounts regarding reduced accuracy of these devices in darker skin, they may worsen healthcare disparities for the Black population. Colvonen et al.¹⁴⁴ suggested that testing of optical wearable devices is often carried out in populations lacking diversity with inaccurate determinations of skin tone and that consumer wearable companies are aware of these flaws but doing little to address them. On the contrary, Bent et al.¹⁴⁵ investigated the reliability of wearable optical heart rate monitors on 53 subjects with a variety of FSTs and showed that that these were unaffected by skin color. However, Colvonen et al.¹⁴⁶ suggested that these results may have been due to small sample size and inaccuracies in Fitzpatrick skin color determination and are unlikely to reflect the true effect of skin color on the accuracy of optical wearable devices. Others also suggest that small sample sizes and inaccurate skin color determination may be leading to misinformation regarding the accuracy of medical wearables in people of color.^134,147