Photochemistry and Photobiology, 2009, 85: 33–44
Review
Application of Nonlinear Optical Microscopy for Imaging Skin†
Kerry M. Hanson and Christopher J. Bardeen*
Department of Chemistry, University of California at Riverside, Riverside, CA
Received 29 July 2008, accepted 5 November 2008, DOI: 10.1111/j.1751-1097.2008.00508.x
(Fig. 1). The epidermis is composed of four stratified layers (in
ascending order: strata basale, spinousum, granulosum, corneum) and typically ranges between 50–1500 lm thick depending upon body site (3). The stratum corneum (SC), the
outermost epidermal layer, is composed of anucleated keratinocytes, and is a biochemically complicated region having an
active role in immunoregulation and barrier homeostasis. Its
main functions are to limit physical trauma and penetration of
topical agents in addition to inhibit extrusion of bodily fluids.
The rest of the epidermis consists of three nucleated layers
below the SC that have live cells in various states of
differentiation as they move toward the surface and form the
SC. These layers host Langerhans cells that are responsible for
regulating immune response in the skin. Melanocytes are also
present in the epidermis, residing primarily in the basale layer
and extending upward into the upper epidermal layers, and
give rise to the degree of skin pigmentation. The dermis lies
below the epidermis and is primarily composed of the
structural proteins collagen and elastin. This layer typically
ranges between 100–500 lm thick and contains hair follicles,
glands and other larger scale structures.
As the brief description above shows, skin is a spatially
heterogeneous system with a large variety of chemical environments. The ability to directly observe both the structural
and chemical properties of skin at different depths would
enhance our understanding of how it functions and what
changes occur in its diseased states. Optical microscopy is the
most commonly used technique for the study of biological
tissue, but until recently, it has been difficult to probe the
structural and chemical properties of the skin using photons
(i.e. optical microscopy) because skin is a strongly scattering
medium that is opaque at ultraviolet and visible wavelengths.
Near-infrared (IR) light can penetrate deeply into the skin, but
most biological molecules and labels do not linearly interact
with such long wavelength photons. The invention of femtosecond near-IR lasers and subsequent demonstration that twophoton fluorescence excitation using near-IR femtosecond
pulses leads to sectioned, 3D imaging of biological cells with
submicron spatial resolution helped usher in a new era in skin
imaging (4).
Nonlinear excitation is achieved by the simultaneous
interaction of two or more IR photons and is usually a v(2)
or v(3) process, and as such nonlinear optical microscopy
(NLOM) is often referred to as multiphoton microscopy
(MPM). A detailed discussion of nonlinear optical processes
ABSTRACT
Recent advances in the use of nonlinear optical microscopy
(NLOM) in skin microscopy are presented. Nonresonant spectroscopies including second harmonic generation, coherent antiStokes Raman and two-photon absorption are described and
applications to problems in skin biology are detailed. These
nonlinear techniques have several advantages over traditional
microscopy methods that rely on one-photon excitation: intrinsic
3D imaging with <1 lm spatial resolution, decreased photodamage to tissue samples and penetration depths up to 1000 lm
with the use of near-infrared lasers. Thanks to these advantages,
nonlinear optical spectroscopy has become a powerful tool to
study the physical and biochemical properties of the skin.
Structural information can be obtained using the response of
endogenous chemical species in the skin, such as collagen or lipids,
indicating that optical biopsy may replace current invasive, timeconsuming traditional histology methods. Insertion of specific
probe molecules into the skin provides the opportunity to monitor
specific biochemical processes such as skin transport, molecular
penetration, barrier homeostasis and ultraviolet radiationinduced reactive oxygen species generation. While the field is
quite new, it seems likely that the use of NLOM to probe structure
and biochemistry of live skin samples will only continue to grow.
INTRODUCTION
An outer layer of skin, roughly 1 mm thick, provides the only
barrier between the body’s soft tissues and the external world.
It is the largest single organ of the body. From a medical
standpoint, diseases of the skin, including cancer, psoriasis,
acne and eczema cost the United States an estimated
$38.6 billion in medical costs every year (1). In addition to
disease, attempts to improve skin quality and counteract aging
processes provide the foundation of the multibillion dollar
cosmesceutical industry (2). Thus, there are powerful medical,
social and commercial motivations for improving our understanding of the structure and function of skin.
Human skin is a complicated multilayer structure that can
be divided into two main regions: the epidermis and dermis
†This invited paper is part of the Series: Applications of Imaging to Biological
and Photobiological Systems.
*Corresponding author email: [email protected] (Christopher J.
Bardeen)
2009 The Authors. Journal Compilation. The American Society of Photobiology 0031-8655/09
33
34
Kerry M. Hanson and Christopher J. Bardeen
microscopy (TPM) is used in these applications. Many
different types of fluorescent molecules have been developed
that are sensitive to a variety of chemical characteristics (pH,
reactive oxygen species [ROS], calcium). In this review, we
provide an overview of recent applications of both types of
imaging to specific problems in skin biology. In addition, we
provide an in-depth summary of recent work from our group
on the measurement of ROS in the skin, and the measurement
of pH levels in response to various types of stimuli.
MATERIALS AND METHODS
Figure 1. A simplified diagram of human skin. Epidermal thickness
depends upon body site being thickest on the palms and soles
(1500 lm) and thinnest around the eyes (10 lm). The stratum
corneum (SC) is the only layer composed of anucleated, terminally
differentiated cells called keratinocytes that are surrounded by a lipidrich extracellular matrix. The SC provides the primary barrier function
of the skin and is be affected by pH. The pH changes dramatically
through the SC where the H+ concentration decreases 100–1000-fold
from pH 5 at the skin surface to pH 7 in the first nucleated layer. All
other epidermal layers contain nucleated keratinocytes. The dermis is
composed primarily of the structural proteins collagen and elastin, but
also contains fibroblasts, glands and hair follicles.
can be found in Ref. (5,6). NLOM has several advantages over
1-photon confocal fluorescence microscopy and optical coherence tomography methods which rely on v(1) processes, i.e.
one-photon absorption or scattering (4,7–11). First, NLOM
permits the use of IR light, which affords greater depth
penetration (<1000 lm) than visible or UV excitation
(<50 lm) (7,8,12). This allows one to obtain physical and
biochemical information on the skin with the submicron
spatial resolution that NLOM affords (7,8,12). Second,
NLOM is confocal-like in that it allows 3D sectioning but
without the use of a pinhole that reduces light collection and
affects image quality (4,7–9,11). Third, because near-IR light is
not resonant with endogenous skin chromophores and the
excitation is localized only in the focal region, photodamage to
the tissue sample is minimized (7,8,12).
In this review, we cover several aspects of the recent use of
NLOM to probe both the structure and chemical characteristics of live skin. We begin with a general description of the
instrumentation and sample preparation methods, followed by
a summary of recently published applications of NLOM for
skin imaging. These applications can be roughly divided into
two main categories. First, one can use endogenous chromophores already in the skin to generate signals such as second
harmonic generation (SHG), coherent anti-Stokes Raman
(CARS) or autofluorescence (AF). These types of imaging rely
on contrast between different tissue compositions and types,
and are used mainly to provide information on structure and
structural integrity within the skin. Second, one can add an
exogenous chromophore to the skin in order to obtain more
specific chemical information. Most of these probe molecules
are present in relatively low concentrations and must be
detected via their fluorescence. Thus, two-photon fluorescence
All NLOM methods use a multiphoton fluorescence microscope, and similar imaging and detection parameters. Ragan
et al. has published a comprehensive review of the instrumentation characteristics required for MPM (8). AF, SHG, CARS
and exogenous probe fluorescence NLOM use a multiphoton
microscope which requires a laser light source, a microscope and
detectors (4,7–12). In a typical experiment, the laser excitation
light travels through the epifluorescence port of the microscope.
A dichroic mirror reflects the laser light and passes the
fluorescence to the detector. Detectors are positioned on
the bottom port of the inverted microscope or the top port of
the upright microscope. The choice of detector depends upon the
type of experiment (fluorescence intensity or lifetime, spectra,
CARS, generalized polarization [GP]) and is discussed in more
detail below. Broadband and bandpass filters placed in the
detection path prevent residual IR from reaching the detectors.
Excitation requirements. In general, nonlinear optical processes can only be driven efficiently using short, intense pulses,
and thus the use of fs or ps laser pulses. As mentioned above,
the ideal wavelength range for skin imaging is the near-IR.
Finally, the laser should be tunable across a wide wavelength
range in order to excite a variety of chromophores. These three
requirements: short pulses, near-IR wavelength and tunability
have led to titanium:sapphire laser oscillators becoming the
most widely used light sources for NLOM. To generate
femtosecond pulses, frequency doubled CW green laser light
pumps a Ti:Sapphire femtosecond laser for excitation between
720 and 900 nm. Most major laser manufacturers offer lasers
which meet these criteria.
Fluorescence and SHG excitation. The use of a tunable
excitation source is important to optimize imaging conditions
in the skin. AF, SHG and exogenous probe fluorescence can be
maximized by selecting the appropriate wavelength (4,7–9,11–
13) (Table 1), and the wide wavelength range of the Ti:Sapphire femtosecond laser oscillator makes it the preferred
excitation source for most types of experiments in the skin.
The epidermis and dermis contain numerous chromophores
that fluoresce (Table 1) upon two-photon excitation between
720 nm and 760 nm. Thus, tuning between this wavelength
range can maximize AF from these molecules. Laser tunability
also allows one to select exogenous fluorescence probes that
can be excited at wavelengths greater than 760 nm in order to
minimize AF, but maximize the probe contribution to the
overall fluorescence signal. It should be noted that in fluorescence experiments, the signal appears wherever the molecular
emission dictates and is not affected by changing the excitation
wavelength, unlike in SHG imaging, where the excitation
wavelength does determine the signal wavelength, since the
Photochemistry and Photobiology, 2009, 85
Table 1. Endogenous skin chromophores.
Chromophore
Fluorescence and SHG
Retinol (20)
NADH (17,97–100)
Vitamin D (20)
Flavins (17)
Melanin (101)
Elastin (101)
Collagen
Fluorescence
SHG (19,101)
CARS
C-H stretch
Sebaceous glands
Adipocytes
Excitation kex (nm)
Emission k (nm)
700–830
340; 690–730
<700
370, 350; 700–730
280–450
300–340; 700–740
450
450–470
450
430
440, 520, 575
420–460
300–340; 700–740
720–960
(tunable range
of TP laser)
Excitation
See (23,29)
420–460
360–480 (kex ⁄ 2)
Emission x (cm)1)
2845 cm)1
2845 cm)1
2956 cm)1
SHG will occur at a wavelength exactly half that of the
excitation wavelength (k ⁄ 2) (14–20). The skin has been found
to contain molecules that exhibit a SHG signal, for example at
collagen interfaces (Table 1). SHG is polarization sensitive, so
a k ⁄ 4 waveplate can be used in the excitation path to achieve
circular polarized light so chromophores like collagen fibers
can be excited uniformly (18,21). As in the case of fluorescence,
the excitation wavelength may be varied in order to maximize
the signal and minimize the background (usually AF). A
narrow-band spectral filter, centered at k ⁄ 2, can be used to
suppress the spectrally broad AF background.
CARS. Coherent anti-Stokes Raman scattering microscopy
requires similar microscope and detector instrumentation as
AF, SHG and exogenous fluorescence. However, it also
requires two laser wavelengths separated by the molecular
vibrational frequency of interest for both excitation of the
sample and detection of the CARS signal. Currently, there are
three modalities to achieve the required two wavelengths: (1)
two narrowband (ps) pulses from a single laser source and an
OPO (22,23), (2) an excitation pulse and a broadband probe
pulse created by white-light continuum (24) or by a fs
Ti:Sapphire (25) and (3) a single, interferometric, shaped pulse
that provides both the excitation and probe pulses (26–28). In
the case of two separate pulse sources, the two beams must be
spatially and temporally overlapped before reaching the
scanning mirrors outside the microscope. In contrast, the use
of a single pulse for both excitation and probe pulses avoids
such a need. Some examples of laser sources used in CARS
microscopy include an Nd:Vanadate laser at 1064 nm and an
OPO at 780–930 nm, and a single or double Ti:Sapphire laser
(730–950 nm) configuration (23,26–30).
Excitation power and sample damage. For two-photon fluorescence, typical average excitation powers are <5 mW for
ex vivo tissues and 50 mW in vivo. Selecting an excitation
power can be a trade-off between image quality (higher power
leads to more signal) and tissue damage (higher power leads to
more photodamage). To a first-order approximation Dunn
et al. showed that two-photon fluorescence signal decreases
exponentially with increasing focal depth; however, it should
be noted that degradation in two-photon-excited signal has
35
many sources including, but not limited to, scattering of
excitation and emission photons, objective N.A., tissue optical
properties and detector choice (31,32). Dunn et al. showed that
with increasing imaging depth comes more scattering of the
excitation photons, which in turn is the primary source for the
reduced image quality seen at greater depths in TPM images
(32). As a result, using higher laser powers increases the
photon density at the point spread function to improve image
quality simply by allowing more photons at the excitation
volume; however, care must be taken to avoid photodamage to
the sample by using too high of an incident laser power (8,32).
Contrast agents such as glycerol, propylene glycol and to a
lesser extent glucose may help with this compromise by
reducing excitation scatter, and improve image contrast and
penetration depth (33). They may remove water through
osmosis and reduce heterogeneity in the index of refraction for
better matching with the NA. Tissue damage is a concern for
higher laser powers. Photodamage in the skin from twophoton excitation can occur three ways: (1) by intracellular
chromophore absorption of radiation, which may be similar to
damage caused by UV radiation, (2) from dielectric breakdown from electromagnetic radiation and (3) from one-photon
absorption of IR (34–36). The dominant mechanism can
depend sensitively on excitation conditions and sample. Of
course, in some cases damage can result from a combination of
factors. An example is the observation of thermal mechanical
damage during TPM optical biopsy of the skin. In this
application, it was found that the majority of photodamage in
skin occurs at the epidermal–dermal junction and results from
the one-photon absorption of IR by melanin causing cavitation (aka explosive evaporation) at the focus (8,32). The
damage was best minimized by reducing the laser repetition
rate to reduce the average energy deposited in the sample, as
opposed to simply reducing the pulse energy, which also
compromised the fluorescence signal. For every NLOM
experiment, conditions must be found that optimize signal
while avoiding sample damage.
Scanning and imaging. There are multiple ways to obtain an
image in NLOM of the skin. High quality images can be
obtained using slow point-by-point scanning. Such methods
scan point-by-point over the x–y plane either by scanning the
excitation point through the use of galvanometer-driver
scanning mirrors (Cambridge Technology, Cambridge, MA)
or by physically moving the sample in the x–y plane through a
motorized stage (H101, Prior Scientific, UK). Typical frame
rates of 0.5 s to 10 s permit the achievement of excellent image
quality (11,37). Unfortunately, such ‘‘slow’’ scanning methods
can be a disadvantage when imaging in the skin because of its
highly heterogeneous environment, which, depending upon the
experiment, can require multiple areas and multiple skin
samples to be imaged to obtain statistically relevant data (38).
Thus, other methods can be used to achieve video rate imaging
(30 frames s)1) and speed up the data collection process. For
example, line-scanning uses a line focus created by a cylindrical
lens in the excitation path to image a full line rather than a
single point (39). Although this method does improve image
acquisition time, it has a disadvantage in that it yields reduced
resolution in the axial direction (point-scanning z-resolution
1 lm; line-scanning resolution >3–6 lm) (31). To simultaneously achieve both rapid scanning and the resolution of
36
Kerry M. Hanson and Christopher J. Bardeen
point-scanning images, multiphoton multifocal microscopy
has proven successful. This method uses an array of lenses that
when rotated focus on multiple spots uniformly over the ximage plane, while a synchronized galvanometric mirror is
used to scan the y-axis (39,40). Bewersdorf et al. showed that
scan rates of 225 frames per second can be achieved; however
the charge-coupled device (CCD) frame rate of 32–67
images s)1 set the limit for the actual image acquisition.
Similar image acquisition speeds can be achieved using a fast,
rotating polygonal mirror, which can scan a line 50 · per
revolution along the x-axis, improving the rate of image
acquisition 100-fold over point-scanning methods. In addition,
resonant mirror (41) and acousto-optic scanners (42) may also
be used to achieve rapid scanning of the laser beam in the x–y
plane. In almost all cases, scanning along the axial (z) direction
is accomplished using a motorized piezo-driven stage to
position the focal spot of the beam at different depths within
the tissue.
In addition to rapid scanning of the laser spot, there are some
specific optical requirements for using NLOM to image skin
samples. When choosing an objective, the index of refraction (n)
of the immersion medium should be closest to the average of the
index of refraction of skin, 1.4 for the least amount of
spherical aberration and the greatest image quality (38).
Because the skin is heterogeneous the n values vary with depth
(stratum corneum n = 1.47, stratum basale n = 1.34 and
dermis n = 1.41), and thus it is impossible to match all n values
of the skin (8,43). Both water and oil high NA (1.3) objectives
have been found to yield high quality images, with oil objectives
collecting more fluorescence when imaging the dermis (44,45).
Typical objectives are 40 · infinity corrected oil or water
objectives (F Fluor, 1.3 NA; Zeiss, F Fluor NA 1.3, Nikon).
Signal detection. After excitation, the sample emits photons
which must be detected in order to generate the final signal.
Depending on the nature of the experiment, one may need to
detect the intensity, spectrum, polarization or temporal profile
of the emitted photons. The detection system must be
interfaced to a computer in order to generate a pixel-by-pixel
image. For NLOM data acquisition for home-built systems,
the SimFCS computer program (Laboratory for Fluorescence
Dynamics, University of California at Irvine, lfd.uci.edu) is a
popular resource available to researchers. We address the
specifications for different NLOM methods used to image skin
in this section. Below we summarize the different detection
modalities in the context of different experiments.
Intensity measurements. Often only simple intensity measurements are needed, which require a single sensitive detector (e.g.
photomultiplier tube [PMT]) or avalanche photodiode for
single-channel collection. Using commercially available detectors, single-photon sensitivity is routinely achievable. The
main challenge is to eliminate unwanted background from the
scattered laser photons or AF. The most common strategy for
suppressing the background photons is to spectrally filter them
out using a high quality bandpass or edge filter.
Spectral measurements. In general, more information can be
obtained if both the intensity and spectrum of the emitted light
is detected. For example, a spectrometer can be placed before a
CCD camera in one channel of the detection pathway to
collect not only intensity data but also emission spectra from
fluorescence or SHG (13,20). If only coarse spectral information is needed, and detected photons are at a premium, then
the emitted light can be divided into two spectral regions and
the intensity in both can be detected simultaneously. For
example, to collect both fluorescence (430–700 nm depending
upon fluorophore) and SHG (370–410 nm, or k ⁄ 2, depending upon k), two detectors for dual channel (a.k.a. dual color)
detection can be employed where a dichroic and filters are
placed in the detection path to separate the two colors.
Another example of this approach is the development of a
trimodal instrument to image both linear (index of refraction
contrast, absorption, scatter) and nonlinear events (twophoton fluorescence, SHG) combining optical coherence
microscopy with spectrally separated SHG and two-photon
fluorescence (46). In this experiment, a dichroic mirror was
used to separate backscattered IR light for the optical
coherence microscopy from the visible light generated by
two-photon fluorescence and SHG. The emission from the
latter two was in turn separated by the placement of an
additional dichroic before two PMT detectors.
In CARS microscopy, where the signal is shifted from the
excitation by an amount equal to the vibrational frequency
within a molecule, spectral resolution of the signal is required.
In the skin, the C–H stretch gives a strong CARS signal at
2845 cm)1. Sebaceous glands (2845 cm)1) and adipocytes
(2956 cm)1) have also been detected (Table 1), and the amide
bands in the structural protein collagen may potentially be
observed as well (23,29,47). These vibrational frequencies
determine the shift of the signal from the excitation wavelength, and as a result, CARS microscopy requires the use of a
red-sensitive detector such as a CCD or PMT when near-IR
lasers are used (23).
Polarization and generalized polarization. Polarization (P) and
GP measurements can be used to monitor molecular orientation and changes in microenvironment, respectively. Polarization measurements detect fluorescence intensity both parallel
(Ipar) and perpendicular (Iperp) to the polarization direction of
the excitation light. In the skin, AF is collected to indicate the
molecular orientation of molecules in the imaged area, with a
high absolute P-value indicating a relatively stationary sample
with little molecular rotation (48,49). P is calculated using
Eq. (1).
P¼
Ipar Iperp
Ipar þ Iperp
ð1Þ
Polarization measurements require the use of a polarizer
before both excitation of the sample as well as the detector
(48).
By combining spectral and polarization resolution, one can
measure the GP. This method is used to identify changes in the
lipid-aqueous microenvironment by measuring the fluorescence from the probe Laurdan, which has been used to probe
the skin barrier (50,51). For example, Laurdan redshifts with
increasing polarity and indicates a more disordered state with
more water molecules present. The GP is calculated using
Eq. (2).
GP ¼
I440nm I490nm
I440nm þ I490nm
ð2Þ
Photochemistry and Photobiology, 2009, 85
where Laurdan fluorescence intensity I is measured at 440 nm
or 490 nm using a dual channel PMT set-up and spatial filters.
Circularly polarized light is used for GP measurements in skin
so all of the randomly oriented molecules are excited equally at
different depths (48).
Fluorescence lifetime imaging microscopy. In the experiments
described above, the time-integrated signal intensity per pixel is
what yields the final image. Variations in intensity provide image
contrast and are usually correlated with structural features or
changes in chemical composition. Intensity measurements are
most useful when one simply wants to detect the presence of
fluorescence that is absent in a control. In cases where the
intensity of the probe fluorescence cannot be directly compared
to a control intensity image due to inhomogeneous labeling (see
pH section below), then fluorescence lifetime imaging microscopy (FLIM) becomes useful. This is because the fluorescence
lifetime of a chromophore is independent of concentration and
inhomogeneities in excitation and emission paths, and thus
FLIM is a powerful method useful for probes whose inhomogeneous distribution in the skin compromises the intensity
images and where a comparative control image cannot be made
with confidence.
There are two ways to collect FLIM data: frequency-domain
or time-domain data acquisition (52, 53). In brief, in frequencydomain FLIM, the fluorescence lifetime is determined by its
different phase relative to a frequency modulated excitation
signal using a fast Fourier transform algorithm. This method
requires a frequency synthesizer phase-locked to the repetition
frequency of the laser to drive an RF power amplifier that
modulates the amplification of the detector photomultiplier at
the master frequency plus an additional cross-correlation
frequency. In contrast, time-domain FLIM directly measures
s using a photon counting PMT and high speed data acquisition
card (52–54).
Skin samples. With approval from a university Internal
Review Board, ex vivo skin can be acquired from what would
be considered waste following plastic surgeries or from patient
volunteers. Human cadaver skin can be obtained from
national skin banks (National Disease Research Interchange,
Philadelphia, PA). Skin equivalents (Epiderm, Epiderm-FT,
MatTek Corp., Ashland, MA) have also been found to be
excellent substitutes for ex vivo and cadaver skin for TPM
studies (43,55–57). Animal subjects have been used too, but
imaging experiments are highly sensitive to movement by a live
subject, so euthanized animals tend to yield the best images
due to their immobility (58).
SKIN IMAGING WITH ENDOGENOUS
CHROMOPHORES
The first area of interest is the use of different NLOM techniques
to image skin in the absence of fluorescent labels. In such ‘‘labelfree’’ experiments, the optical response is dominated by the
properties of the endogenous chromophores, many of which are
listed in Table 1 (17,19,20,23,29,59–63). To date this type of
imaging has been used primarily for examining micron scale
structures to differentiate between normal and diseased skin,
usually for clinical diagnostic purposes. More recent reports
have surfaced that couple microscopy with spectroscopy to
37
identify not only the location of fluorescence but also what
species is giving rise to that fluorescence (13,47). There are three
types of endogenous signals researchers commonly use to
diagnose the health of the skin: AF, SHG and CARS. Below we
describe applications of these detection modalities to some
specific skin biology problems in more detail.
Skin dermatopathology and biopsy
The detection of fluorescence from endogenous chromophores
by NLOM is being developed as a replacement for traditional
biopsy methods that require physical removal of the tissue
sample followed by a histological analysis (7,64). Development
of a real-time, noninvasive microscopy technique to distinguish between healthy and diseased tissue could offer a
significant improvement over the current highly-invasive and
time-consuming biopsy method available.
AF from keratin and NADH, along with collagen SHG, can
differentiate between normal, precancerous and squamous cell
carcinoma tissue in hamster cheeks (65). Similarly, Lin et al.
distinguished between basal cell carcinoma (BCC) and normal
tissue on cross-sectioned, formalin-fixed tissues using AF and
SHG (21). The ratio between AF and SHG signals (MFSI) was
significantly greater for BCC tissues relative to normal dermis
(MFSIcancer clumps = 0.9; MFSInormal = 0.35), which may
show that optical biopsy of skin cancers is close-at-hand. AF
and SHG have also been used to investigate dermal diseases
in vivo that are associated with changes in collagen and elastin
such as scleroderma or graft versus host disease (66). In addition,
such noninvasive methodology may lead the way to improve our
understanding of wound healing or matrix destruction of
invasive tumors within the skin and in vivo (66). For example,
Navarro et al. successfully used TPM to follow wound-healing
in full-thickness guinea pig skin wounds (58).
In addition to using the AF and SHG signal intensities to
identify diseased states, resolving the temporal and spectral
characteristics of the fluorescence can also be used for
diagnostic purposes (33,67). For example, by combining a
spectrometer postexcitation and before a CCD detector,
spectral imaging of the skin can be used to spectrally resolve
different autofluorescent chromophores while concomitantly
identifying their cellular location in an image for identification
of diseased areas in a bulk skin sample (47). However, as Paoli
et al. showed, more information is needed on the fluorescence
characteristics of skin tumors by comparing them with
traditional histopathology results before noninvasive fluorescence biopsies can become a reality (68).
Another promising technique for the evaluation of skin
pathology is CARS microscopy. The contrast mechanism is
based on the response of molecular vibrations, and can
differentiate between structures in the skin such as the sebaceous
glands and adipocytes, as well as lipid distributions in the skin,
like the lipid-rich extracellular matrix in the SC (23). It has the
potential to detect DNA and protein vibrational signals, and
potentially collagen, all of which could be highly useful to
differentiate between normal and diseased skin (47).
Skin aging
Although not technically a disease, skin aging is still of great
concern because of the concomitant reduction in barrier
38
Kerry M. Hanson and Christopher J. Bardeen
function, as well as the personal-care desire for ‘‘wrinklefree’’ skin. Similar to a pot-hole in a road with a poor
foundation, wrinkles in the skin appear due to the breakdown of collagen and elastin in the dermis, but the
mechanisms by which deterioration occurs are not fully
understood. NLOM using AF and SHG signals is a
promising approach to increase our understanding of the
pathways that lead to skin aging.
For example, Lin et al. found that the SHG signal
decreased and AF increased with age on excised formalinfixed facial skin (18). Koehler et al. found identical results
in vivo using the forearm of human subjects (66). The results
from both studies are consistent with the solar ultravioletinduced progressive damage of matrix metalloproteases
digestion of dermal collagen and elastosis which increases
truncated elastotic (and autofluorescent) material as one
ages. Both groups assign a SHG-to-AF-aging-index-of-dermis,
or SAAID, ratio to quantify the degree of aging of the sample.
The lower the SAAID indicates the greater degree of dermal
damage. The SAAID ratio in particular may prove a valuable
parameter for the skin care industry. Currently, there are little, if
any, concrete values used to evaluate the appearance of the skin.
With the SAAID parameter, one can envision a tool that
monitors the progress of drugs upon skin appearance in vivo,
quickly and noninvasively.
In addition to looking at the end-result of aged skin,
NLOM can be used to study how aging occurs. For example, it
is possible to look more closely at the mechanism by which
collagen deteriorates by monitoring collagen SHG signal
(69,70). Sun et al. showed that denaturation along a fibril
occurs at a faster rate than the rate of global fibril denaturation within a bundle, indicating that the forces that keep the
fibrils bundled together are stronger than those that keep the
individual collagen molecules aligned in a single fibril (69).
They induced denaturation thermally, heating their collagenrich rat tail tendon samples between 40C and 70C. From
these results, one could envision, with further research,
therapeutics that can target the denaturation processes that
lead to skin wrinkling as we age.
SKIN IMAGING WITH EXOGENOUS
CHROMOPHORES
Information about the structural integrity and health of skin
obtained from NLOM using endogenous chromophores can
be further refined by the use of exogenous chromophores (71).
The use of labeling molecules to improve contrast has a long
history in histology, and fluorescent labels can provide useful
information about structures that cannot be distinguished
otherwise. In addition to imaging static structures, exogenous
fluorescence probes have proven to be the only effective means
to obtain more specific biochemical information from the skin,
although CARS microscopy may prove to be the exception to
this rule in the future (23,29). To study the biochemical
environment of the skin, exogenous chromophores such as
organic dyes and inorganic quantum dots can be applied to
the skin surface, incubated for a period of time and then the
skin is imaged (55–57,71–74). When exogenous chromophores
are used, their concentrations are typically much lower than
that of the endogenous chromophores, and fluorescence is the
only practical detection modality. Thus, in most of the
applications described below, TPM is the method used for
imaging.
Dermatopathology and skin biopsy
The AF and SHG experiments described in the previous
section tend to be relatively nonspecific at the molecular level.
Exogenous chromophores, on the other hand, can be used to
selectively label chemical components within the skin. Fluorescein antibodies and rhodamine lectins have been used to
label nerves, blood vessels and hair follicles in the skin (71).
TPM has also been used to study cell migration and
colocalization of Langerhans (immunoregulatory) cells in the
skin around the nucleus, and enhanced green fluorescent
protein-labeled dermal cells in mice (75). In this study, the
authors also used the intrinsic SHG of dermal collagen to act
as a reference point to the migrating cells. Each study
exemplifies the advantages of TPM over traditional histopathology or confocal laser methods. Histopathology is labor
intensive and slow, while one-photon methods do not provide
the depth penetration needed to actually get information at the
different epidermal depths. In addition, NLOM has the
advantage that it can image in real-time, which is crucial for
cell-tracing studies used to follow pathogenesis (75). TPM
methods may one day lead to pathological diagnoses in vivo to
complement AF and SHG biopsy, if fluorescent probes can be
approved for human subject use.
Transport through the stratum corneum
The SC is remarkable in its ability to inhibit the penetration of
topical agents including sunscreens and industrial chemicals.
However, in some cases, one desires to increase the permeability of the SC for the purposes of drug delivery. In either
case, a better understanding of the barrier function of the SC is
required. The SC is the thin (on average 10 lm) brick-andmortar-like layer of anucleated protein-rich keratinocytes
surrounded by a lipid-rich extracellular matrix (Fig. 1). How
to impede or enhance transport through the SC is an area of
active research.
TPM has been used to directly observe the transport of
nanoparticles, over-the-counter crèmes, and hydrophobic
and hydrophilic probes across the SC (38,76–80). The latter
TPM experiments on hydrophobic rhodamine B hexyl ether
and hydrophilic sulforhodamine B shed light on the complicated and unique intracorneocyte and extracellular routes
that both hydrophobic and hydrophilic molecules can take
through the SC (76–78). Several experiments have been used
to correlate transport rates with structural properties of the
SC. To understand the diffusion paths available in the SC,
Sun et al. looked at the microenvironments within the SC
(48). They coupled TPM with P and GP detection to study
oleic acid (OA)-induced structural changes in the extracellular lipid matrix. Polarization images showed that
application of OA leads to a decrease in P, indicating an
increase in dynamic disorder and reduced lipid packing.
More specifically, GP images showed that application of OA
leads to a decrease in GP, regardless of the polarization of
the light, and that GP varies between sites. This indicates
that water most likely infuses into microdomains as the OA
fluidizes the lipid multilamellae in the extracellular matrix,
Photochemistry and Photobiology, 2009, 85
8.0
5
150
Figure 2. The fluorescence lifetime-sensitive pH probe 2¢,7¢-bis(2-carboxyethyl)-5-(and-6)-carboxyfluorescein (BCECF). Because
BCECF’s fluorescence lifetime (sf) changes with pH, which is not
affected by inhomogenous labeling, it can be used to accurately
monitor pH in the skin. At pH 4.5, sf = 2.75 ns, and at pH 7.1
sf = 3.97 ns.
4.0
and subsequently improves transport of some molecules
through the extracellular SC matrix (81). Further research
has shown that transport can be improved not only through
the extracellular matrix but also through the keratinocytes
themselves. Lee et al. used TPM to monitor the effects of
the depilatory (hair-removal) agent (thioglycolate) upon
transport through the SC, and found a ‘‘bubbling effect’’
within the keratinocytes in response to thioglycolate, essentially making small pores within the cells (82). They also
found that in some cases keratinocytes completely detached
from the skin surface leaving behind a ‘‘halo’’ of intracellular lipid matrix. Traditional methods such as staining or
histology could not identify these effects within the keratinocytes themselves. In fact, these experiments are a perfect
example of how NLOM can be coupled with traditional
staining and histology methods to get information that the
latter simply cannot detect. The authors coupled the TPM
image data with data acquired from traditional staining
experiments that looked at the structure of the lipid matrix,
and could make the final determination that thioglycolate
improves transport of model drugs through both the
intracellular and extracellular spaces for at least 24 h after
application In addition to these intensity and polarization
measurements, Bird et al. showed that two-photon FLIM
can detect changes in corneocyte AF lifetime following
application of the hormone ethinyl estrodiol in ethanol.
They found that the intracorneocyte AF lifetime decreased
postdrug application, showing that this method may be
useful to monitor drug delivery pathways in the skin as well
(83).
Although at the beginning stages, these three methods
(intensity, polarization and lifetime) are markedly advancing
our understanding of how drugs are transported through the
skin. Further research in this field may help provide simple,
cost-effective drug delivery systems that also have fewer sideeffects.
39
Figure 3. Fluorescence intensity of the lifetime-sensitive pH-probe
BCECF and corresponding pH maps of mouse skin at different
epidermal depths. The pH maps were calculated using BCECF’s
lifetime values, and not its intensity. The intensity images show the
importance of taking lifetime measurements. BCECF’s fluorescence
intensity is greatest at high pH, but the areas in the skin with the
greatest intensity are in reality at low pH. The intensity variations are
due to inhomogeneous labeling by the fluorophore (72,84,85).
Characterization of the pH of the stratum corneum
Transport through the SC barrier is determined by the
barrier’s chemical properties. Determining the chemical composition of the SC is also relevant for understanding how this
barrier can be repaired when damaged by common ailments
such as eczema or diaper rash or by more severe wounds such
as burns. There is evidence that the SC barrier function is
greatly influenced by the pH gradient present in the SC
(Fig. 1). Thus, a first step in the study of the biology of the SC
is to measure its pH at different SC depths and under different
conditions (43,55,84,85).
To measure pH in the SC, FLIM measurements of skin
incubated with the lifetime-sensitive pH probe 2¢,7¢-bis-(2-carboxyethyl)-5-(and-6)-carboxyfluorescein (BCECF) (Fig. 2) have
been conducted. Both the probe intensity and fluorescence
lifetime depend on local pH. Exogenous chromophores have the
benefit that they are able to provide information about chemical
processes occurring in the skin, but caution must be used when
interpreting intensity images in this case. In particular, BCECF
inhomogeneously labels the skin such that probe intensity in a
pixel may appear more or less depending simply upon the
amount of label present in that pixel. If one used a simple
intensity measurement to detect local variations in skin pH, one
might erroneously assume that a region having higher intensity
would have a higher pH, as opposed to having a greater local
concentration of dye at the same pH. For this experiment, one
must use FLIM to obtain reliable data. The lifetime (sf) of
BCECF changes with pH—at pH 4.5, sf = 2.75 ns and at
pH 7.1 sf = 3.97 ns (55,81). FLIM measurements identified the
presence of 1 lm diameter acidic microdomains in the lipidrich extracellular matrix compared to the neutral intracellular
space of the corneocytes (Fig. 3) (55). The changing ratio of
acidic microdomains to neutral regions is the source of
the change in pH over the short SC distance. The images
in Fig. 3 exemplify the importance of using FLIM when
labeling is heterogeneous. In a homogeneous environment, the
fluorescence intensity of BCECF is greater at neutral pH than at
acidic pH (86). However, as Fig. 3 shows, BCECF does not label
uniformly, and rather its intensity is greatest in areas of acidic
pH. Thus, although the intensity images show a bright fluorescence in some regions, the corresponding pH is not neutral, but
rather is acidic. Previously, bulk methods were employed to
determine pH as a function of SC depth, where skin layers were
successively tape-stripped and the pH was measured with a pH
probe (87,88). In contrast, two-photon FLIM allowed the pH to
40
Kerry M. Hanson and Christopher J. Bardeen
be characterized with submicron spatial resolution at different
depths and without disrupting the sample.
Further work found that the formation of acidic microdomains occurs at the stratum granulosum-SC interface and is
regulated by the sodium-proton exchanger NHE1 (76). In
addition, the acidic SC surface is not fully developed at birth,
and rather acidic microdomains at the SC-SG interface
develop postnatal (85). Niesner et al. also used TP FLIM on
artificial skin constructs and found that an identical pH
gradient exists to that found in mammalian skin, which could
further research on barrier function without the need for
human or animal tissues (43).
Detection of reactive oxygen species in the skin
In addition to detecting the presence of protons, other small
molecule species like ROS can also be monitored by using
appropriate probe molecules. ROS are highly reactive derivatives of oxygen and include superoxide anion, hydroxyl
radical and singlet oxygen. They are formed naturally during
cellular respiration, and through energy transfer to or reaction
with O2 following UVB (280–320 nm) and UVA (320–450 nm)
absorption by skin chromophores including urocanic acid,
NADH, riboflavin and melanin (89–94). The photogenerated
ROS have been implicated in photoaging and possibly skin
cancer, since overexpression of ROS leads to oxidative stress
which can induce photoaging, immunomodulation, DNA
damage and actinic keratosis (skin cancer precursors)
(95–98). Thus, it is of interest to measure the presence of
ROS in different skin layers and see whether their level can be
suppressed or enhanced under different experimental conditions. Our research has explored the use of TPM to study the
effects of solar UV radiation on the generation of ROS in the
skin (55–57,99).
ROS in skin can be detected by exogenous chromophores
including dihydrorhodamine (DHR) (Fig. 4). DHR is nonfluorescent until reaction with ROS when it becomes fluorescent
rhodamine-123 (R123, kem = 535 nm). The reaction scheme is
given in Fig. 5. By simply measuring the increase in R123
fluorescence, we can estimate how many ROS are generated by
solar irradiation. The challenge in monitoring the level of ROS
in living skin tissue is to accurately image the fluorescent R123
molecules that indicate the presence of ROS. The lower
Figure 4. Fluorescence intensity images of skin (z = 30 lm) incubated with DHR before (a) and after (b) UVB irradiation. The
fluorescence in (a) results from AF and DHR conversion to R123 due
to mitochondrial respiration. The increase in fluorescence in (b) results
from R123 that forms from the reaction of DHR with ROS. R123
fluorescence is detected primarily in the cytoplasm of the keratinocytes,
which may result from inhomogeneous labeling by DHR.
Figure 5. The ROS probe dihydrorhodamine (DHR) is nonfluorescent
until it reacts with ROS to form fluorescent rhodamine-123 (R123).
DHR is not a selective reactant and may react with many other ROS
than those listed above. It also does not localize in nuclei or in cell
membranes and cannot identify if ROS are generated in these regions
on keratinocytes. Other ROS probes may prove to be useful to provide
more data on these cellular locations.
scattering and greater depth penetration of 800 nm laser pulses
allowed us to accurately measure the formation of ROS at the
subcellular level in live skin tissue. Using these methods, we
found that UVB irradiation (equivalent to 2 h noonday
summer sun in North America) of ex vivo skin samples
generates 14.7 mmols of ROS in the SC and 0.01 mmols in all
of the viable layers for the average adult-size face of 258 cm)2
(55). Because DHR may not have labeled cell membranes,
nuclei and other cellular components, these experiments may
underestimate the level of ROS that are truly generated;
however, they do show that ROS are generated in significant
amounts by a UVB dose often obtained on a summer day (55).
Armed with a protocol to quantify the amount of ROS in
live skin, we can now examine how the generation of ROS
can be modulated by external factors. A simple way to
minimize the number of ROS generated is by the application
of FDA-approved UV-filters used in sunscreens (57). Using
TPM, we can look below the surface layer of SC and applied
sunscreen to see what effects these molecules have on the cells
of the epidermis. Our measurements showed that octocrylene
(OC), octylmethoxycinnamate (OMC) and benzophenone-3
(B3) (Fig. 6) all reduced the number of ROS generated in the
epidermis following irradiation by solar-simulated UVB-UVA
if they remained on the skin surface (Fig. 7). However, as the
skin was incubated for t = 20 or t = 60 min with OC, OMC
and B3 formulations, the UV-filters penetrated below the SC
surface. These molecules then absorbed the solar-simulated
Figure 6. Three FDA-approved UV-filters commonly used in overthe-counter sunscreens: (a) benzophenone-3 (B3), (b) octocrylene (OC)
and (c) octylmethoxycinnamate (OMC).
Photochemistry and Photobiology, 2009, 85
Figure 7. R123 fluorescence intensity of epidermis (z = 60 lm) after
20 mJ cm)2 UVB-UVA radiation. Skin applied with crème containing
B3, OC or OMC and incubated t = 0 min show a decrease in
fluorescence compared to the placebo. After t = 60 min, the fluorescence of B3, OC or OMC-applied skin is greater than the placebo
fluorescence. Identical results were found for all nucleated epidermal
layers at t = 60 min, indicating that the UV-filters penetrated the skin
surface and generated ROS themselves (57).
UVB-UVA (20 mJ cm)2 [10 min. summer sun in North
America]) and generated ROS deep within the nucleated
layers of the epidermis (Fig. 7). These results show that if
OC, OMC and B3 penetrate the skin surface they can
generate more ROS in the nucleated epidermis than if
sunscreen wasn’t used; however, a concomitant attenuation
of UV at the skin surface (i.e. from reapplication of the
sunscreen) should inhibit OC, OMC and B3 from sensitizing
ROS because no UV light could reach them to initialize the
ROS sensitization.
The vehicle (crème) plays a significant role in the degree
of penetration of a UV-filter or any topical agent. Ideally,
one should formulate a vehicle to improve retention of a
UV-filter on the skin surface, so that it acts in a manner
similar to latex-paint (57). In addition, topically applied
antioxidants have been found to reduce ROS levels in the
nucleated epidermis, although typically a large amount must
be present in the formulation to significantly reduce the
number of UVB-UVA-induced ROS (57). TPM has also
been used to show that dietary lutein reduces UV-induced
ROS in mouse epidermis (100). These results illustrate that
TPM can provide more detailed data on the efficacy of a
sun protection product. Sunscreens do an excellent job at
protecting against sunburn when used correctly. However,
sunburn may not be the only risk factor for skin cancer, and
reactions indistinguishable to the naked eye, such as those
instigated by ROS, may play a significant role as well.
Clearly, there appears to be room for more research in
photoprotection science.
SUMMARY
The research summarized in this paper has shown that
NLOM not only improves upon traditional skin research
methods, but also provides new information on skin properties and biochemistry. NLOM is proving to be a promising
clinical tool, especially by providing an alternative to tradi-
41
tional biopsy and histology methods. AF and SHG of the skin
have differentiated between normal, diseased and cancerous
tissues, and followed wound healing or matrix destruction of
invasive tumors, all of which may lead to in vivo optical
biopsy measurements. With further advances, multiphoton
endoscopes that can image noninvasively in vivo may become
even more common (101). Because TPM requires labeling of
the skin, it may not be a strong candidate for in vivo biopsy;
however, it is showing to be highly applicable for ex vivo
dermatopathology, which may one day amend or replace
traditional ex vivo histology. Perhaps more importantly, TPM
is also providing basic scientific information on barrier
properties and biochemistry of the skin that until now has
been impossible to obtain in unfixed tissues. Based upon these
studies, NLOM is likely to play a key role in answering
important questions in skin biology. Examples, to name just a
few, include: the role of calcium in barrier homeostasis, how
rosacea develops, how hormones affect hair loss, and the role
of ROS in skin cancer. The future of this microscopy
technique for advancing skin studies appears to be very
bright.
Acknowledgements Work was supported by the National Science
Foundation, grant MCB-0344719.
DISCLOSURES
KMH has consulted in the sunscreen industry.
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