UCLA
UCLA Previously Published Works
Title
Rapid and accurate estimation of blood saturation, melanin content, and epidermis
thickness from spectral diffuse reflectance.
Permalink
https://escholarship.org/uc/item/2tx5r31d
Journal
Applied optics, 49(10)
ISSN
1539-4522
Authors
Yudovsky, Dmitry
Pilon, Laurent
Publication Date
2010-04-01
DOI
10.1364/AO.49.001707
Peer reviewed
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University of California
Rapid and Accurate Estimation of Blood Saturation,
Melanin Content, and Epidermis Thickness from
Spectral Diffuse Reflectance
Dmitry Yudovsky, and Laurent Pilon
§
University of California, Los Angeles
Henri Samueli School of Engineering and Applied Science
Mechanical and Aerospace Engineering Department
Biomedical Inter-Department Program
Los Angeles, CA 90095-1597
§
Corresponding Author
Phone: +1 (310)-206-5598, Fax: +1 (310)-206-4830
E-mail: pilon@seas.ucla.edu
December 15, 2010
ABSTRACT
This paper presents a method to determine chromophore concentrations, blood saturation,
and epidermal thickness of human skin from diffuse reflectance spectra. Human skin was
approximated as a plane-parallel slab of variable thickness supported by a semi-infinite layer
corresponding to the epidermis and dermis, respectively. The absorption coefficient was
modeled as a function of melanin content for the epidermis and blood content and oxygen
saturation for the dermis. The scattering coefficient and refractive index of each layer were
found in the literature. Diffuse reflectance spectra between 490 and 620 nm were generated
using Monte Carlo simulations for a wide range of melanosome volume fraction, epidermal
thickness, blood volume, and oxygen saturation. Then, an inverse method was developed to
retrieve these physiologically meaningful parameters from the simulated diffuse reflectance
spectra of skin. A previously developed accurate and efficient semi-empirical model for
diffuse reflectance of two layered media was used instead of time-consuming Monte Carlo
simulations. All parameters could be estimated with relative root mean squared error less
than 5% for (i) melanosome volume fraction ranging from 1 to 8%, (ii) epidermal thickness
from 20 to 150 µm, (iii) oxygen saturation from 25 to 100%, (iv) blood volume from 1.2
to 10%, and (v) tissue scattering coefficient typical of human skin in the visible part of the
1
spectrum. Similar approach could be extended to other two-layer absorbing and scattering
system.
OCIS codes: 100.3190: Inverse problems, 170.1470: Blood or tissue constituent monitor-
ing, 170.1870: Dermatology, 170.3660: Light propagation in tissues, 170.3880: Medical and
biological imaging, 170.6510: Spectroscopy, tissue diagnostics
1 INTRODUCTION
Diffuse reflectance spectroscopy has found many applications in non-invasive monitoring of
biological tissues [1–7]. This technique investigates tissue structure, chromophore concen-
tration, and health by measuring the tissue’s optical properties. Commercially available
devices typically analyze experimental data using the modified Beer-Lambert’s law to deter-
mine the relative concentrations of tissue chromophores such as melanin, blood, water, or
hemoglobin in arbitrary units [6, 8–13]. However the tissue scattering coefficient cannot be
retrieved [10,14]. Alternatively, diffuse reflectance data processed with the diffusion approxi-
mation [15] can yield absolute chromophore concentration and measure the tissue’s scattering
coefficient [16] which is related to tissue microstructure [16–19]. However, this technique re-
quires emitter-detector separation up to 1 cm thus limiting the spatial resolution of these
devices [20, 21].
Current spectroscopic techniques are based on the assumption that tissue is homogeneous
and that properties are independent of depth [6,7,9,11,22,23]. In reality, most bodily organs
such as skin, the intestine, or the cervix are protected by a thin lining called the epithelial
layer [24]. While the organ is typically composed of connective tissues and perfused with
blood vessels and nerves, the protective epithelial layer is bloodless and consists of structured
cell layers [24]. Differences in cellular structure and chemical composition give rise to distinct
optical properties making the assumption of tissue homogeneity questionable [25]. In skin,
for example, the outer epidermal layer is pigmented by melanin which absorbs strongly in
the UV while the inner dermal layer is pigmented by blood which absorbs in the visible and
near-infrared parts of the spectrum [26]. Furthermore, the thickness of the epidermal layer
may vary with anatomical location, gender, and age [27–29].
Multi-layer optical models of tissue have been developed [13,30–32] to study the effects of
tissue structure and chromophore distribution on light propagation. However, such mo dels
are computationally intensive and cannot be used in real-time clinical applications [33]. Semi-
empirical models of light transfer have been developed to accelerate computation without
significant loss of accuracy [34–38]. Mantis and Zonios [34], for example, developed a semi-
empirical model for diffuse reflectance of two-layer media. They used their model in an
inverse method to determine the optical properties of two-layer tissue phantoms of variable
epithelial thickness. However, their method requires that the bottom layer be scattering
but non-absorbing [34] which is not the case for most organs [24, 26]. In addition, Tsumura
et al. [36] developed a semi-empirical model for diffuse reflectance of two-layer scattering
and absorbing media based on the modified Beer-Lambert’s law. The authors assumed that
the thickness of the top layer was equal to 70 µm [36]. Recently, Yudovsky and Pilon [35]
developed a semi-empirical model that predicts the diffuse reflectance of strongly scattering
two-layer media. The model accounted for (i) absorption and anisotropic scattering in both
2
layers, (ii) variable thickness of the top layer, and (iii) internal reflection at the medium/air
interface. It was shown to accurately predict the diffuse reflectance of skin [35].
The objective of this study is to develop an inverse method based on our semi-empirical
model [35] and to assess its robustness in estimating the scattering coefficients, chromophore
concentrations in both epidermis and dermis, and the epidermal thickness from spectral
diffuse reflectance of human skin.
2 BACKGROUND
Human Skin
Skin is the largest organ of the human body representing a total surface area of approximately
1.8 m
2
and a total weight of approximately 11 kg for adults [39]. The epidermis and dermis
are the two main layers. They are separated by the basement membrane and rest on the
subcutaneous fat layer [39]. The topmost layer of the epidermis is called the stratum corneum
and is composed of dead cells embedded in a lipid matrix. The rest of the epidermis is
mainly composed of keratinocytes, melanocytes, and langerhans [39]. Melanocytes synthesize
melanin, the skin protein mainly responsible for skin color. Melanin is contained in organelles
know as melanosomes which are distributed throughout the epidermis [39]. Depending on
genetic factors and UV light exposure, melanosomes occupy 1 to 43% of the epidermal volume
corresponding to lightly or darkly pigmented skin, respectively [40–42]. Epidermal thickness
varies with b odily location and ranges between 20 and 150 µm [26, 42–44].
The dermis, located beneath the epidermis, is responsible for the skin’s pliability, me-
chanical resistance and temperature control. It contains touch, pressure, and temperature
receptors as well as sebaceous and sweat glands and hair follicles [39]. The dermis is com-
posed of collagen fibers perfused by nerves, capillaries, and blood vessels [26, 45, 46]. The
thickness of the dermis ranges between 450 and 650 µm [27,47]. Depending on body location
and tissue health, the volume of blood in the dermis ranges between 0.2% and 7% [42,48,49].
Approximately half of the blood volume is occupied by erythrocytes (red blood cells) which
are responsible for oxygen transfer from the lungs to the rest of the body [42, 49]. Ery-
throcytes are composed mainly of hemoglobin molecules which reversibly bind to oxygen
molecules in the lungs to form oxyhemoglobin. Hemoglobin is known as deoxyhemoglobin
once it has released its oxygen molecules. The ratio of oxyhemoglobin molecules to the total
number of hemoglobin molecules in the blood is the so-called oxygen saturation denoted by
SO
2
. Hemoglobin absorption dominates the total absorption of the dermis in the visible
range [26, 39, 46]. Furthermore, the spectral extinction coefficient of oxyhemoglobin differs
significantly from that of deoxyhemoglobin. Thus, the color of the dermis depends on the
average oxygen saturation of its blood content.
Skin Properties Measurement
Various techniques exist to measure chromophore concentrations and blood saturation of
human skin. Commercially available non-invasive, optical devices typically measure these
quantities in a small region 1 to 2 cm in diameter and report a device specific melanin
3
(MI) and erythema (EI) index [40, 50]. The MI corresponds qualitatively to the darkness of
skin while EI corresponds to the redness or inflammation of skin. Such devices have been
used to predict the risk of melanoma skin cancer [40] and as dosimetry feedback during
laser treatment of port-wine stains [51] and acne [52]. Recently, hyperspectral imaging in
the visible and near-infrared parts of the spectrum has been used to determine the spatial
distribution of oxygen saturation in the human skin [12]. This technique has been applied
clinically to study diabetic neuropathy [53] and predict the healing potential of diabetic
foot ulcers [4, 54]. Such devices typically assume a homogeneous tissue structure and do
not model changes in the scattering coefficient with wavelength, biological state, or from
patient to patient. Thus, only relative chromophore concentration in arbitrary units can
be reported [10, 12, 13]. Furthermore, epidermal thickness and blood volume cannot be
determined [10].
Epidermal thickness varies naturally with age, gender, and body location [27–29, 55]. It
may also increase or decrease due to external stimuli. For example, UV exposure of human
skin has been shown to increase the thickness of the epidermis in addition to increasing its
melanin content [56,57]. On the other hand, smoking has been shown to decrease epidermal
thickness [55]. Epidermal thickness can be measured reliably with punch biopsy whereby a
sample of the skin is removed and analyzed ex vivo [27,55,57]. This invasive technique can
be painful and destroys the sample. Alternatively, non-invasive measurements of epidermal
thickness can be made with techniques such as optical coherent tomograph or ultrasound
[58,59]. However, these techniques are primarily sensitive to the tissue’s scattering coefficient
therefore simultaneous determination of chromophore concentration is difficult [60, 61].
The Radiative Transfer Equation
Biological tissues such as skin are generally absorbing and strongly scattering media [33].
Light transfer through such turbid media is governed by the radiative transfer equation
(RTE) written as [15]
ˆs · ∇I(ˆr,ˆs, λ) = −µ
a
(λ)I(ˆr,ˆs, λ) − µ
s
(λ)I(ˆr,ˆs, λ) +
µ
s
(λ)
4π
Z
4π
I(ˆr,ˆs
i
, λ)Φ(ˆs
i
, ˆs, λ)dΩ
i
(1)
where I(ˆr,ˆs, λ) is the spectral intensity at location ˆr in a unit solid angle dΩ around direction
ˆs expressed in W/cm
2
·sr·nm. The linear spectral absorption and scattering coefficients are
denoted by µ
a
(λ) and µ
s
(λ), respectively and are expressed in cm
−1
while the scattering
phase function is denoted by Φ(ˆs, ˆs
i
, λ). The Henyey-Greenstein scattering phase function
is an approximate expression that accounts for the anisotropic nature of scattering and is
given by [62],
Φ(ˆs
i
, ˆs, λ) =
1 − g(λ)
[1 + g(λ)
2
− 2g(λ) cos Θ]
3/2
(2)
where Θ is the angle between ˆs and ˆs
i
and g(λ) is the Henyey-Greenstein asymmetry factor
used extensively in tissue optics [33,63,64]. The values of g(λ) measured for the epidermis and
dermis were approximately the same and ranges between 0.73 and 0.82 in the visible range
[46]. In order to account for the magnitude and anisotropy of the scattering phenomenon,
4
