Ex vivo and in vivo second-harmonic-generation
imaging of dermal collagen fiber in skin:comparison of imaging characteristics
between mode-locked Cr:forsterite
and Ti:sapphire lasers
Takeshi Yasui,* Yu Takahashi, Masahiro Ito, Shuichiro Fukushima, and Tsutomu Araki
Graduate School of Engineering Science, Osaka University, 1-3 Machikaneyama-cho,
Toyonaka, Osaka 560-8531, Japan*Corresponding author: t‑yasui@me.es.osaka‑u.ac.jp
Received 2 September 2008; revised 22 December 2008; accepted 26 December 2008;
posted 5 January 2009 (Doc. ID 100978); published 26 January 2009

Second-harmonic-generation (SHG) microscopy is an interesting new tool for observing dermal collagenfiber in skin. However, conventional SHG microscopy using a mode-locked Ti:sapphire laser suffers fromlow penetration depth and a slow image acquisition rate caused by scattering and absorption in tissue,
making it difficult to use for in vivo applications on human skin. We develop an SHG microscope equippedwith a mode-locked Cr:forsterite laser with a long wavelength and compare its imaging characteristics
with that of a Ti:sapphire-laser-based SHG microscope for the measurement of dermal collagen fiber inanimal and human skins. The results indicate the suitability of the Cr:forsterite laser-based SHG
microscope for in vivo imaging of human skin. © 2009 Optical Society of America
OCIS codes:
170.1870, 170.3880, 170.6960, 180.4315, 190.4160.

1. Introduction

Dermal collagen fiber plays an important role in determining morphological and mechanical propertiesof skin as a structural protein. Therefore, there is agreat need for an appropriate monitoring tool thatcan reveal the three-dimensional (3D) structure ofdermal collagen fibers in vivo. Recently, secondharmonic-generation (SHG) microscopy with anear-infrared femtosecond pulse laser has received
interest as an optical probe method to investigate collagen fiber in biological tissues due to the fact that
biological SHG light is specifically generated by collagen molecules [1,2]. This SHG light also provides
unique imaging modality: high contrast, high spatial
resolution, optical 3D sectioning, noninvasive, anddeep penetration. By using endogenous SHG as a0003-6935/09/100D88-08$15.00/0© 2009 Optical Society of America
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contrast mechanism, this microscopy enables visualization of the structures of collagen fiber in tissues
without additional staining and hence avoids phototoxicity and photobleaching. By taking advantage ofthese characteristics, SHG microscopy has been
applied to visualize detailed structures of dermal collagen fiber in skin [3,4]. In addition, combining thistechnique with polarization methods, namely, polarization-resolved SHG microscopy, can give additional
insights with regard to the orientation of collagenfiber in tissues and has been used to quantify thecollagen fiber orientation in the dermis [5–7].
In most previous studies, a mode-locked
Ti:sapphire (Ti:S) laser with a center wavelengtharound 800 nm was used as the light source to generate SHG light around 400 nm [2–7]. However,scattering and absorption of the skin at these wavelengths often limits in vivo monitoring of dermal collagen fiber in a reflection configuration. Multiple

scattering events greatly reduce the number of ballistic photons of the incident laser light, which isthe dominant source of nonlinear excitation in deep
focusing nonlinear microscopy [8], although the scattering does not significantly degrade the profile ofthe point spread function at the focal spot [8,9]. In
the return path, 400 nm SHG light is largely attenuated by severe multiple scattering and absorption.
These effects lead to an insufficient penetration
depth for in vivo monitoring of the dermal collagenfiber across the stratum corneum and epidermis
layers in a reflection setup. For polarization-resolved
SHG microscopy, multiple scattering events further
depolarize the detected SHG light, resulting indecreased sensitivity to collagen fiber orientationin deep regions of tissue [7]. Another problem within vivo applications for human skin is the risk of biological photodamage in the vicinity of the focal spot
caused by nonlinear absorption and/or ionization of
the incident laser light [10,11]. The problems of conventional Ti:S laser-based SHG microscopy can beaddressed by the use of an ultrashort pulse laser
with a longer wavelength to suppress multiple scattering and absorption in tissue. In a promising alternative laser source for SHG microscopy, the Ti:S
laser is replaced with a mode-locked Cr:forsterite(Cr:F) laser centered around 1250 nm, which takes
advantage of the fact that absorption and scatteringin tissue is minimum near a wavelength of 1300 nm[12,13]. Recently, combined SHG and third-harmonic-generation microscopy using a Cr:F laser was effectively applied to in vivo developmental biology ofzebrafish embryos [14], the optical biopsy ofa hamster oral cavity [15], fixed human skin [16],and teeth [17]. Few studies have directly compared
the SHG imaging performance of Ti:S and Cr:F
lasers, although plant tissue cell damage underirradiation of tightly focused beams has been compared [18]. Here we report the construction of
two equivalent reflection-mode SHG microscopes
equipped, respectively, with a Ti:S laser and a Cr:
F laser. We compare their imaging characteristics
by ex vivo and in vivo measurements of dermal
collagen fiber in animal and human skins.
2. Experimental Setup

As a light source for the SHG microscopes, weemployed an 800 nm Ti:S laser (Spectra Physics,
MaiTai; 100 fs pulse duration, 800 mW average power,
80 MHz repetition rate) and a 1250 nm Cr:F laser(Avesta Project, Moscow Region, Russian Federation,
CrF-65P; 135 fs pulse duration, 200 mW average
power, 73 MHz repetition rate). The experimental setup for the SHG microscope is illustrated in Fig. 1. For
rapid acquisition of SHG images, the laser beam isscanned in two dimensions by a pair of galvano
mirrors (GM). After passing through relay lenses(RL1 and RL2), the laser beam is focused onto thesample with an objective lens (OL; Nikon Instruments, Melville, New York, CFI Plan 50 × H,
50 magnification, 0:9 numerical aperture, 350 μm

Fig. 1. Experimental setup: M, mirror; HS, harmonic separator;
ND, neutral density filter; GM, galvano mirrors; RL1 and RL2,
relay lenses; OL, oil-immersion objective lens; PZT, piezoelectrictransducer; F, infrared-cut filter; PMT, photon-counting photomultiplier tube; L, lens.

working distance, oil immersion). The immersion oil
for the OL decreases the reflection loss and sphericalaberration caused by mismatching the refractive index at the sample surface. The OL can be moved along
the optical axis by a piezoelectric transducer (PZT,
stroke length of 350 μm ). Combining the GM and
the PZT enables optical 3D sectioning images of
SHG light to be taken. The average power of bothlasers on the sample was set to 40 mW. Although
the dwell time of the high power can be reduced by adjusting the excitation power with depth, we kept thelaser power at this power level for ease of comparison.
Whereas the SHG light induced in the sample propagates mainly in the same direction as the incident
laser light, a portion of the generated SHG light isbackscattered into the sample. This backscattered
SHG light, which is the main signal source of SHG
imaging in the reflection setup, is collected via the
OL. The SHG light is descanned by the GM and separated from the laser light by a harmonic separator(HS, reflected wavelength of 400 nm for Ti:S SHG lightor 625 nm for Cr:F SHG light) and an infrared-cut
filter (F). The SHG light is detected by a combined
photon-counting photomultiplier tube (PMT) andpulse counter. To suppress the wavelength dependence of the PMT sensitivity, we used a different
PMT with the same sensitivity at wavelengthsof 400 and 625 nm for each SHG microscope(Hamamatsu Photonics K.K. 7155-20 with a countsensitivity of 2:7 × 105 s−1 pW−1 for Ti:S SHG lightand H8259-01 for Cr:F SHG light with a count sensitivity of 2:3 × 105 s−1 pW−1 ). Since the PMT for the Ti:S
SHG microscope has an internal prescaler that scales
the signal down by a factor of 4, we multiply the measured photon count by four. Since the field of view ofthe microscope can be changed by scanning the angleof the GM, we adjusted the field of view size to fit observed structures of dermal collagen fibers within therange of the image.
In this way, we designed a descanned collection
system for SHG light; many nonlinear microscopes are based on a nondescanned collection
system for efficient collection of scattered SHG light.
Our interest is in the in vivo application ofpolarization-resolved SHG microscopy for examining
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skin, and since multiple scattering depolarizes the
SHG light and hence decreases the sensitivity tocollagen fiber orientation [7], it is important to rejectunnecessary multiple-scattered SHG light efficiently when probing the deep regions of skin withpolarization-resolved SHG microscopy. To reject multiple-scattered SHG light, a descanned collection system is more suitable than a nondescanned systembecause the former detects the ballistically backscattered SHG light with little depolarization rather
than multiple-scattered SHG light.
3. Results
A. Ex Vivo Observation of Animal Skin

We first took ex vivo SHG images of a block sample
excised from porcine skin, scanning the observation
plane parallel to the skin surface along the depth
direction, namely depth-resolved SHG imaging. We
used the back skin of a Yucatan micropig purchasedfrom Charles Rivers Laboratories (Wilmington,
Massachusetts) as a sample. The excised skin tissue
was embedded in optical cutting temperature (OCT)compound (Sakura Finetek, Tokyo, Japan) and snap
frozen in liquid nitrogen. A frozen block of the skintissue was sliced to 5 mm thickness parallel to theskin surface. Then the dermal collagen fiber wasmeasured across the stratum corneum andepidermis layers by the two SHG microscopes.
Figures 2(a) and 2(b) show two series of depthresolved SHG images measured with the Ti:S SHG
and Cr:F SHG microscopes with the probing depth
from the skin surface varied in 50 μm intervals. To
compare two SHG images with different photon
counts, we scaled them to grayscale images with
256 gradations for each image. Note however thatthe imaging area on a sample for the Ti:S SHG microscope is not consistent with that for the Cr:F SHG
microscope. Although it might be better to performthe image comparison with the same area on a sample, it is difficult for the two SHG microscopes to image exactly the same area on a sample because thetwo laser sources used in this study were set in different experimental rooms and we had to construct
two independent SHG microscopes. Consecutive
depth-resolved SHG images at depth intervals of
10 μm are also shown as a movie (Media 1) of Fig. 2.
The area of each image is 400 μm by 400 μm, composed of 256 pixels by 256 pixels. The image acquisition rate is 10 s=image. In the Ti:S SHG images, theSHG signal first appears at a depth of 80 μm below
the skin surface [see the movie (Media 1) in Fig. 2],
which indicates that the analytical spot enters thecollagen-rich dermis layer from the no-collagen
epidermis layer. The distribution of the collagen fiberin the papillary dermis was imaged. However,
the Ti:S SHG image rapidly became out of focus
when the probing depth exceeded 110 μm, and the
SHG light almost disappeared beyond a depth of
200 μm. Because of this, the measurable depth of
the Ti:S SHG microscope was limited to a shallow
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Fig. 2. Depth-resolved SHG imaging of dermal collagen fiber in
porcine skin measured with (a) Ti:S SHG microscope and (b) Cr:FSHG microscope at 50 μm intervals. The image size is 400 μm by
400 μm and the image acquisition rate is 10 s=image. The movie(Media 1) shows consecutive change of their depth-resolved
SHG images at 10 μm intervals.

region of the dermis. The Cr:F SHG microscope
allowed depth-resolved SHG imaging to a deeper
region of the dermis. As for the Ti:S SHG, the Cr:FSHG signal from dermal collagen fiber starts to appear after a depth of 80 μm. A structural change inthe dermal collagen fiber is clearly visualized, depending on the probing depth. Within a depth region
from 90 to 110 μm, the fine collagen fiber is distributed uniformly while the capillary vessels and pores
without collagen content appear as black spots,
which are specific characteristics in the papillary
dermis layer. Even beyond a 120 μm depth, at which
point the Ti:S SHG microscope rapidly loses its spatial resolution accompanied by a decrease of signalto-noise ratio, the Cr:F SHG microscope maintains asufficient spatial resolution to image the collagenfiber in the reticular dermis layer, in which the dermal collagen fibers gradually become thick. The Cr:FSHG image starts to become out of focus and lose
sharpness beyond a 250 μm depth, however, a SHG
signal was detected even at a depth of 350 μm.
Since the maximum probing depth of the present
Cr:F SHG microscope is limited by the working distance of the OL and the PZT ð¼ 350 μmÞ, using a longer working distance will increase the penetration
depth beyond 350 μm. In this way, a comparison of
depth-resolved SHG images of porcine skin for the
Ti:S SHG and Cr:F SHG microscopes indicates thatthe latter is superior to the former with respect tohigh penetration power, which is important forin vivo observation of dermal collagen fiber in humanskin. In usual nonlinear microscopes, sphericalaberration of the OL, caused by mismatching of
the refractive index at the sample surface, distorts
and broadens the point spread function profile atdeep probing regions, resulting in an out-of-focus
SHG image and/or a weak SHG signal at the deep
regions of the skin [19]. Refractive indices of the skintissues are 1.51 for stratum corneum, 1.34 for epidermis, and 1.41 for dermis [20]. Since the refractive
index of the immersion oil used here is 1.515, themismatch of the refractive index at the skin surface(oil–corneum interface) is 0.005. Furthermore, when
focusing the laser light on the layered tissue structure of the skin by the OL, additional mismatch of
the refractive index occurs at interfaces between different tissue layers: 0.17 for the corneum–epidermis
interface and 0.105 for the epidermis–dermis layer.
These mismatches result in spherical aberration.
Multiple scattering in tissues attenuates both theincident laser light and the detected SHG light. In
Fig. 2, the resolution of the Ti:S SHG images is comparable to that of the Cr:F SHG images for depths
between 50 and 100 μm. However, the resolutionof the former degrades rapidly beyond a depth of
100 μm. Since the wavelength dependence of thespherical aberration is considered to be small, thedifference in the image degradation between the
two SHG microscopes is mainly due to the wavelength dependence of multiple scattering. The SHG
image degradation at deep regions of the skin canbe simply suppressed by introducing a confocal
pinhole in the descanned collection system, at theexpense of the SHG signal. Furthermore, both the
SHG efficiency and the image quality improve if
the spherical aberration of the OL is well compensated by adaptive optics [21,22].
We next evaluated the imaging characteristics ofboth SHG microscopes for rapid measurement. The
probing depth was fixed at 140 μm from the surfaceof the porcine skin, so that SHG images of the dermal
collagen fiber were measured. Figures 3(a) and 3(b)show Ti:S SHG images at an acquisition time of 10
and 1 s, respectively. The area of each image is
200 μm by 200 μm, composed of 256 pixels by 256
pixels. A sufficient photon count was not obtainedfor either 1 or 10 s acquisition times. Therefore, it
is difficult to measure a fine SHG image rapidly atthis probing depth with the Ti:S SHG microscope.
Cr:F SHG images for measurement times of 10
and 1 s are shown in Figs. 3(c) and 3(d). Comparing
the two Cr:F SHG images shows no degradation ofimage quality for a shorter image acquisition time
because the photon count of the 10 s SHG image
greatly exceeds the 256 gradations of the grayscale
image. A clear SHG image of the dermal collagen fiber was obtained at the measurement time of 1 s.
Furthermore, the photon counting rate achieved by
the Cr:F SHG was several times higher than that
achieved by the Ti:S SHG, despite the fact that thepulse width of the Cr:F laser is longer than that of
the Ti:S laser. These results demonstrate the potential of Cr:F SHG imaging for in vivo applications onhuman skin from the viewpoint of a fast imaging rate.
The concentration gradient of the dermal collagenfiber along the depth direction is another importantfactor determining the morphological properties ofFig. 3. Rapid SHG imaging of dermal collagen fiber in porcine
skin. Ti:S SHG images at acquisition times of (a) 10 s and(b) 1 s, and Cr:F SHG images at acquisition times of (c) 10 s and(d) 1 s. The image size is 200 μm by 200 μm. The probing depth
is fixed at 140 μm from the surface of the skin.
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skin. Cross-sectional imaging by SHG light canreveal the gradient of the collagen fiber concentration in skin. Recently, SHG OCT was achieved by
combining SHG imaging and OCT techniques andapplied for cross-sectional imaging of collagen fiberin tissues [23,24]. However, the time required to acquire one cross-sectional image is longer than 5 min,even though coherent detection is used to enhance
the weak SHG light. SHG microscopy also enables
cross-sectional SHG imaging, namely, SHG tomography, by B-mode scanning by a tightly focused beam
spot [25]. The high penetration capability and
SHG efficiency achieved by the Cr:F SHG microscope
are attractive for such SHG tomography of dermal
collagen fiber. We performed SHG tomography ofthe back skin of a hairless mouse using the Cr:FSHG microscope. The preparation protocol of theskin specimen is similar to that used for the porcine
skin. Figure 4(a) shows a series of cross-sectional
SHG images measured with the Cr:F SHG microscope with the position of the B-mode scan laterally
varied at 45 μm intervals. Confocal cross-sectional
images of the reflectance fundamental light werealso measured simultaneously, as shown in Fig. 4(b).
Consecutive cross-sectional SHG and confocal reflectance images at 7:5 μm lateral intervals are also illustrated as a movie (Media 2) of Fig. 4. The area of
each image is 300 μm width by 300 μm depth, composed of 256 pixels by 256 pixels. The imaging rateof the SHG tomography was 10 s=image, which ismuch smaller than that achieved by SHG OCT despite the fact that the image quality of our SHG tomography is comparable to that of SHG OCT. This isbecause photon counting detection was applied forthe weak SHG light in addition to the high penetration power and SHG efficiency for the Cr:F laser. A
comparison of the SHG images and confocal reflectance images reveals a line-shaped signal along
the epidermis surface in the SHG images. Since
the epidermis does not contain collagen fiber, weconsider that this signal is due to third-harmonicgeneration light caused by local optical inhomogeneities at the surface [16]. While no SHG area wasobserved underneath the skin surface, indicatingthat the epidermis layer does not contain collagen fiber, a strong SHG signal appeared in the dermis
layer. Although it is difficult to evaluate a detailed
cross section of the dermal collagen fiber due to insufficient depth resolution, the epidermal–dermal
junctions, epidermal prolongations, and pores wereresolved. In addition, the concentration gradient of
the dermal collagen fiber in the skin could be estimated from the cross-sectional SHG image, although
we have to consider the decrease of the SHG signal
depending on the probing depth, caused by multiple
scattering and absorption in tissue. The effect of
SHG signal attenuation depending on depth is compensated by ratio imaging of SHG light to confocalreflectance light.
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Fig. 4. Cross-sectional images of dermal collagen fiber in hairlessmouse skin measured at lateral intervals of 45 μm: (a) Cr:F SHG
images and (b) confocal reflectance image. The image size is
300 μm width by 300 μm depth, and the image acquisition rate
is 10 s=image. The movie (Media 2) shows a consecutive changeof their cross-sectional images at 7:5 μm lateral intervals.

B. In Vivo Observation of Human Skin

When we apply SHG imaging to in vivo measurements of human skin, the risk of biological photodamage caused by nonlinear absorption and/or
ionization in the vicinity of the focal spot must be
considered. Although fluorescent antibody experiments have found that DNA damage of the epidermis
caused by a 60 mW Ti:S laser irradiation is comparable to that caused by slight sunburn by 0.6-MED UV
irradiation [26], such photodamage should not be ignored [10,11]. On the other hand, microspectroscopic
comparison of the two lasers shows that the use of a
Cr:F laser greatly reduces the cell damage to planttissues [18]. Furthermore, little damage was observed on zebrafish embryos [14], hamster oral mucosa [15], and even human skin [27] by long-term
irradiation by a tightly focused Cr:F laser beam.
Based on these previous studies on biological photodamage and our comparisons of the imaging characteristics of the two lasers, we can conclude that the
Cr:F SHG microscope is the best choice for in vivo
observations of dermal collagen fiber in human skin.
An interesting application for in vivo observationsof dermal collagen fiber is the assessment of cutaneous photoaging, which is necessary in fields suchas anti-aging dermatology and skin cosmetic development. Biochemical, histochemical, and immunohistochemical experiments show that cutaneous
photoaging is closely related to the amount andstructure of dermal collagen fiber in the skin [28],and hence in vivo monitoring of dermal collagen fiberwith the Cr:F SHG microscope may contribute
greatly to the assessment of cutaneous photoaging.
As a preliminary experiment for in vivo assessmentof cutaneous photoaging, we here demonstrate invivo depth-resolved SHG imaging of dermal collagenfiber of the forearm skin of a slightly sunburned 57year-old male and a well sunburned 37-year-old maleby use of the Cr:F SHG microscope. The inside andoutside of the forearm skin are selected as in vivo
measurement points to investigate the difference
in dermal collagen fiber with and without cutaneous
photoaging. Figures 5(a) and 5(b) show a series ofdepth-resolved SHG images of the inside and outsideof the slightly sunburned forearm (57-year-old male)when the probing depth is varied at 50 μm intervals
from the skin surface. Consecutive depth-resolved
SHG images at 10 μm depth intervals are also
illustrated as movies (Media 3) in Fig. 5. The areaof each image is 600 μm by 600 μm with 256 pixels
by 256 pixels. The image acquisition rate is
2 s=image. A detailed change in the collagen fiber distribution along the depth was visualized with highcontrast images, although the movie (Media 3) ofFig. 5 exhibits fluctuations due to insufficient fastening of the forearm. On the inside of the forearm skin,
the SHG images start to appear from a shallow skindepth and show that the fine collagen fiber of thepapillary dermis layer is populated at shallow regions while the thick collagen fiber of the reticulardermis layer is observed at deep regions. The images

Fig. 5. In vivo depth-resolved Cr:F SHG imaging of forearm skinof a slightly sunburned 57-year-old male measured at 50 μm intervals (a) inside and (b) outside the forearm skin. The image size is
600 μm by 600 μm, and the image acquisition rate is 2 s=image. The
movie (Media 3) shows consecutive change of their depth-resolved
SHG images at 10 μm intervals.
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also visualize pores, capillaries, and skin groove. On
the outside of the forearm skin, the depth at which
the SHG signal starts to appear is slightly deeper
than for the inside forearm skin, indicating somesmall thickening of the epidermis caused by slight
photoaging. Also, the characteristics of the dermal
collagen fiber on the inside forearm skin shown in
Fig. 5(a), namely, fine collagen fiber at shallow regions and thick collagen fiber at deep regions, werereproduced on the outside of the forearm skin. Thatis, there are no appreciable differences in the dermal
collagen fiber between the inside and the outside of
the slightly sunburned forearm. We next performedin vivo depth-resolved SHG imaging of the dermal
collagen fiber on the inside and outside skin of thewell sunburned forearm (37-year-old male), as shownin Figs. 6(a) and 6(b) and the corresponding movie(Media 4) (600 μm by 600 μm image size with 256 pixels by 256 pixels, image acquisition rate of 2s=image,and depth interval of 50 μm for the figure and 10 μm
for the movie). On the inside of the forearm skin, theSHG images of the well sunburned forearm were similar to those of the slightly sunburned forearm [seeFig. 5(a)], indicating little influence of cutaneous
photoaging on the inside of the forearm skin. Conversely, on the outside of the forearm skin, the depth atwhich the SHG signal starts to appear on the wellsunburned forearm is deeper than that on theslightly sunburned forearm, indicating greater thickening of the epidermis and/or loss of dermal collagenfiber caused by advanced photoaging. Furthermore,
the collagen fiber structure varies more graduallyalong the depth direction than for the outside skinof the slightly sunburned forearm. These results
reflect the difference in dermal collagen fiber
between photoaged skin and nonphotoaged skin.
For a more detailed evaluation of cutaneous photoaging, a combination of SHG imaging and two-photon
autofluorescence imaging [29,30] or polarizationresolved SHG microscopy for measuring collagenfiber orientation [31,32] will be useful.
4. Conclusions

We developed two equivalent reflection-mode SHG
microscopes equipped with mode-locked Cr:F and
Ti:S lasers for ex vivo and in vivo imaging of dermal
collagen fiber in skin and compared their SHG imaging characteristics. For skin penetration capability,
the Cr:F SHG microscope achieved a potential for
penetration to a depth of over 350 μm, twice as deepas achieved by the Ti:S SHG microscope. Furthermore, comparison of the image acquisition rate between the two SHG microscopes indicated thesuitability of the Cr:F SHG microscope. We performed in vivo depth-resolved SHG imaging of human forearm skin with and without sunburn using
the Cr:F SHG microscope. This imaging clarified
the difference in the dermal collagen fiber between
photoaged skin and nonphotoaged skin. The resultsof this study show that Cr:F SHG microscopy has the
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Fig. 6. In vivo depth-resolved Cr:F SHG imaging of forearm skinof a well sunburned 37-year-old male measured at 50 μm intervals(a) inside and (b) outside the forearm skin. The image size is
600 μm by 600 μm, and the image acquisition rate is 2 s=image.
The movie (Media 4) shows consecutive change of their depthresolved SHG images at 10 μm intervals.

potential to become a powerful tool for in vivo clinicalapplications on human skin.
This work was supported by Grants-in-Aid for
Scientific Research Nos. 19650117 and 20240044
from the Ministry of Education, Culture, Sports,
Science, and Technology of Japan. We also thankthe Cosmetology Research Foundation, Japan, forfinancial support.
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