In vivo multiphoton tomography with a wavelength-tunable femtosecond laser has been performed to investigate the autofluorescence intensity of major endogenous fluorophores of human skin in dependence on the excitation wavelength. In high-resolution multiphoton images of different skin layers, clear trends were found for fluorophores like keratin, NAD(P)H, melanin as well as for the elastin and collagen networks. The analysis of the measurements is supplemented by additional measurements of fluorescence lifetime imaging and signal-decay curves by time-correlated single-photon counting.
©2010 Optical Society of America
In vivo multiphoton tomography allows detailed and high-resolution imaging and provides unique insight into structure and function of living tissue [1–6]. It has become a promising method for a variety of applications like skin cancer diagnosis [7,8], determination of skin aging [9,10], tissue engineering  and in vivo drug monitoring [12,13]. With multiphoton excitation several intrinsic fluorophores inside the skin can be readily used as biomarkers of native structure and even cell metabolism without the need of exogenous contrast agents. These fluorophores such as collagen, elastin, melanin, flavin adenine dinucleotide (FAD), reduced nicotinamide adenine dinucleotide (NADH) and porphyrins are not equally distributed inside the skin but vary in abundance within the different layers: the epidermis, the dermis, and the subcutaneous tissue. In order from outermost (surface) to innermost (deepest) layer, the epidermis consists of the stratum corneum, the stratum granulosum, the stratum spinosum and the stratum basale, the farthest epidermal layer from the skin surface. The dermis, located below the epidermal dermal junction, is mainly composed of structural proteins such as collagen and elastin structures. Overlapping emission spectra of the fluorophores, however, make it difficult to clearly separate the individual contributions. One way of distinction is to spectrally resolve autofluorescence and taking advantage of the different spectral emission maxima to identify the individual fluorophores [14,15]. Most studies concerning the identification of different skin fluorophores so far have been performed with fixed excitation wavelength; however, a further separation can be achieved by varying the excitation wavelength inside the NIR spectral range. Utilizing this so-called “optical window of tissue”  with small absorption and scattering coefficients allows exciting different endogenous fluorophores with varying efficiency down to a depth of 200 μm . Many fluorophores have been studied in vitro and in vivo and their absorption spectra for one- and two-photon excitations are well known [2,4,6,14,17–19]. With this knowledge at least the major contributions to the fluorescence can be assigned to specific fluorophores. In addition to the autofluorescence, second harmonic generation (SHG) can occur, a nonlinear optical process, in which two photons interacting with a nonlinear material form a photon with twice the energy of the initial photons. Inside the skin in particular collagen structures can produce SHG light [20,21]. SHG light is instantaneously generated with a highly unsymmetrical directionality mainly parallel to the direction of the incident light due to the phase matching condition. Nevertheless, the SHG light looses the directionality after multiple scattering events in deeper tissue and can therefore also be detected in reflection geometries. Generally, the very low efficiency of nonlinear excitation processes makes it necessary to use pulsed excitation lasers providing high peak intensities. The time-averaged two-photon excited fluorescence and SHG signals are proportional to P2/τ , with P being the mean power and τ the pulse duration.
All measurements reported in this paper were carried out with the CE-certified clinical multiphoton tomograph DermaInspect (JenLab GmbH) [4,6], a laser class 1M device, which has already been used in a variety of in studies for in vivo non-invasive imaging of human skin including skin cancer investigations [7,8], in vivo drug screening of cosmetic and pharmaceutical compounds , diffusion of nanoparticles into the skin [22,23] as well as for basic research. The DermaInspect can even be equipped with a module for fluorescence lifetime imaging (FLIM), i.e. to measure pixelwise signal-decay times . In previous publications in which the DermaInspect was employed for skin imaging usually one or two excitation wavelengths were chosen and kept fixed, e.g. λ = 750 nm for the excitation of autofluorescence and λ = 820 nm for SHG imaging, respectively . However, the signal dependencies on different excitation wavelengths for in vivo imaging have not been addressed so far. In this paper, we present different aspects of in vivo optical sectioning of human skin like the wavelength dependence of the emission intensity of the fluorophores as well as the separation of SHG and autofluorescence components of collagen signals. Variation of the excitation wavelength allows identifying main fluorophores by comparison with dependencies known from literature and reveals how the different fluorophores can be optimally excited and detected with the DermaInspect system. The results are also intended to further clarify and facilitate the use of multiphoton imaging for in vivo clinical use. In addition to the multiphoton images FLIM has been performed and is presented here to exemplarily illustrate SHG and autofluorescence contributions in the signal decay.
Multiphoton tomography was carried out with the multiphoton tomograph DermaInspect (JenLab GmbH) customary equipped with a femtosecond titanium:sapphire laser as excitation source (Mai Tai XF, Spectra Physics). The laser which is tunable in the wavelength range 710 - 920 nm generates ~100 fs pulses with a repetition rate of 80 MHz. It was equipped with a prism-based chirp-compensation unit (DeepSee module, Spectra Physics) to pre-compensate for the group-velocity dispersion of the optics. The DeepSee allows varying the pulse duration at the sample position. The DermaInspect was further furnished with a flexible mirror extension arm housing the focusing optics for easy positioning of the optics onto the skin  and a one-channel module for FLIM measurements based on time-correlated single-photon counting with a temporal resolution of 250 ps. Imaging was achieved by focusing the laser pulses onto the skin with high-NA oil-immersion focusing optics. The same focusing optics was used to collect the non-linear optical response signals, which were generated inside the femto-liter focal volume. The signal light was separated from the excitation light with a dichroic beamsplitter and detected with a photomultiplier tube (PMT, Hamamatsu 7732). The DermaInspect provides a lateral resolution of 0.4 - 0.6 μm and an axial resolution of 1.2 −2.0 μm . In one scan, a maximum area of about 350 x 350 μm2 can be imaged with typically 512 x 512 pixels. For multiphoton tomography the axial position of the scan area inside the skin can be changed by z-positioning of the focusing optics down to a depth of about 200 μm. Actual images are obtained by encoding the pixelwise-collected signal intensities into 8-bit-grey values. Usually, a color-glass filter (BG 39) was used to block residual back-reflected light of the excitation laser to enhance the image contrast and to protect the PMT from being damaged by excessive light. The autofluorescence and SHG-signal components were in some cases additionally spectrally selected with further bandpass filters (BP 395/11 and BP 460/60, AHF). Time-correlated single-photon counting was employed to measure signal-decay curves at each pixel for FLIM. For the FLIM images the resolution was reduced to 128 x 128 pixels to reduce the measurement time and minimize movement artifacts between measurements. Typically, about six minutes were necessary for one series of images for excitation wavelengths from λ = 720 nm to λ = 880 nm resulting from a measurements time of about 13 s per frame and the necessary time for tuning of the laser wavelength and the prism positions.
Changing the excitation wavelength for ultra short pulses is in general accompanied by a change of the pulse duration at the sample position due to the wavelength dependence of the group velocity dispersion of the material. In order to be exclusively sensitive to the wavelength dependence of the signal the pulse duration at the sample was measured and minimized using the DeepSee unit and the mean power P was adjusted for each wavelength λ such that the ratio P2/(A*τ) was kept constant. The wavelength dependence of the area of the focal spot A was estimated to be proportional to λ2. Mean powers corresponding to 10 mW at 780 nm were used down to a depth of 50 μm; for larger depths 20 mW and 30 mW for the data shown in Fig. 7 and 9 . The measurements of the pulse durations were performed using a commercial autocorrelator (CARPE, APE) with the delay-line module placed into the beam path directly behind the laser. The autocorrelation signal was measured directly in the focal plane of the focusing optics and the pulse durations were deduced from fitting the autocorrelation traces with Gaussians. The DeepSee unit was used to minimize the pulse durations behind the focusing optics by carefully adjusting the amount of negative dispersion to compensate the wavelength dependent group velocity dispersion induced by the optics.
For the measurement of the transmission of the signal-beam path of the DermaInspect light of a pulsed xenon-gas discharge lamp was directed along the pathway of the actual signals, i.e. from the sample position backwards through the focusing optics and the flexible mirror arm. The spectrum of the transmitted light was recorded at the position of the detector and corrected for the emission characteristics of the lamp.
All multiphoton measurements were performed on the left forearms of two female volunteers. Obviously, individual differences and different sample locations naturally lead to variations in the depths of the epidermal layers, nevertheless the layers itself could be clearly identified. In preparation of the measurements, the skin was cleaned and disinfected. At the position selected for imaging a drop of water was applied and a metal ring, holding a microscope slip was fixed onto the skin with double adhesive tape. The metal ring provided a “fixed” connection between the skin and the focusing optics of the DermaInspect to prevent losing the position onto the skin between single measurements. Nevertheless, in vivo studies are generally affected by motion artifacts, which are always present in living human subjects. Relatively large wavelength steps were chosen for the scans to minimize the total measurement time while covering at the same time a relative large wavelength range, because of the difficulty for human subjects to stay steady over an extended period.
3. Results and discussion
3.1 Characteristics of the multiphoton tomograph
Essential for the clear interpretation of measured signals is the knowledge about the transmission properties of the signal pathway and the signal-detection efficiency of the detector. Figure 1 shows the transmission of the signal pathway with and without the color-glass filter BG 39 in place. The use of the filter reduces the transmission in particular at the long wavelength side of the curve, which is necessary to prevent excitation stray light from reaching the detector. With the filter in place (dotted line, Fig. 1) the detection beam path transmits mainly between 350 - 650 nm with a maximum around 500 nm. The dotted line indicates the sensitivity of the PMT, which has a maximum around 365 nm and goes to zero below 300 nm and above 700 nm. Figure 1 exhibits the normalized relative transmittance (maximum is normalized to 100%). Absolute transmission values of about 20% and 5.5% were obtained at 651 nm and 665 nm, respectively, by the use of diode lasers. These numbers represent an upper limit of the absolute signal transmissions. The overall signal sensitivity results from convolution of spectral transmittance and detector efficiency and is maximal in the range 420-490 nm.
Figure 2 summarizes pulse durations derived from autocorrelation measurements in the sample plane of the focusing optics. The figure shows the pulse durations with and without the commercially available “low dispersion option” (LD option, Spectra Physics) which was used to obtain the multiphoton images presented in this paper. Higher than linear order dispersion terms are not compensated for by the DeepSee unit leading to longer pulses than the pulses originally emitted by the laser.
3.2 Optical sectioning at different excitation wavelengths
Multiphoton images of different layers, the stratum corneum, the stratum spinosum, the stratum papillare, the stratum basale and from the epidermal-dermal junction have been obtained for different excitation wavelengths. In Fig. 3 (left), an in vivo multiphoton image of the skin surface (stratum corneum) with an excitation wavelength of λ = 720 nm is depicted showing detailed structure with distinguishable hexagonal shapes. The image brightness primarily arises from autofluorescence light of keratin, which is the main fluorophore in the stratum corneum . With increasing excitation wavelength, the signal intensity decreases as can be seen from Fig. 3 (right) which shows a series of multiphoton images with excitation wavelengths in the range λ = 720 nm to λ = 880 nm in steps of 20 nm. A similar decrease of intensity with increasing excitation wavelength is visible in Fig. 4 which shows a multiphoton image recorded at a depth of 22 μm with an excitation wavelength of λ = 720 nm (Fig. 4 left) and the same wavelength series (Fig. 4 right). The images show cells of the stratum spinosum. In this skin layer, NADH and NAD(P)H are the major fluorophores [6,13]. Owing to the fluorescence of NAD(P)H even the mitochondria inside the cells become nicely visible. Typically, the nuclei show much less fluorescence than the cellular cytoplasm and appear as dark spots. The image intensity significantly drops and is close to zero for excitation wavelengths above λ = 800 nm. This spectral dependence can be understood from the characteristics of the one-photon absorption spectra. The coenzyme NADH exhibits maxima at 290 nm and 351 nm, NADPH a maximum at 336 nm for one-photon absorption, respectively. Above 400 nm, the one-photon absorption of both coenzymes goes to zero. Two-photon-absorption spectra can exhibit broader peaks than one-photon absorption spectra due to different selection rules, nevertheless, the main features of the one-photon absorption spectra can approximately be applied to the two-photon absorption spectra. Hence, above 800 nm NAD(P)H can hardly be excited by two-photon absorption, however, depending on intensity the less efficient three-photon excitation can take place.
The coenzyme FAD, another important fluorophore for skin imaging, with a one-photon absorption maximum around λ = 460 nm would cause a signal raise above λ = 800 nm, but does not seem to significantly contribute to the autofluorescence in the imaged layer. Figure 5 shows a multiphoton image with an excitation wavelength of λ = 720 nm and (as in Fig. 4 right) a wavelength series in the range λ = 720 nm to λ = 880 nm measured at a depth of 55 μm inside the stratum papillare of the dermis. In this layer, a different spectral dependence of the image intensity on the excitation wavelength is observed. In the series of images shown in Fig. 5 (right) some regions become darker while other regions become brighter with increasing wavelength. These signals can be almost completely assigned to SHG produced the interaction of the excitation light with collagen structures . Collagen shows both, autofluorescence and SHG light, but under the condition of multiphoton excitation, the autofluorescence is much weaker than the SHG intensity. This can be confirmed by separating the different signal contributions in decay-time measurements or spectrally using accurate filters, which transmit or block the SHG light, respectively. However, the efficiency of the SHG from collagen structures depends on the excitation wavelength . Chen et al. reported for ex vivo human-skin measurements a maximum of SHG for an excitation wavelength of 800 nm . Palero et al. reported an oscillating dependence of the SHG intensity of collagen fibers on the laser wavelength , which had also been suggested by Zuomi et al. , and found that the optimal excitation wavelength for producing SHG light from collagen fibers is about 837 nm. In addition to the SHG efficiency, also the detection efficiency has to be taken into account. The apparent signal raise in Fig. 5 with increasing wavelength, even above 800 nm, results mainly from the higher transmission/detection efficiency of the DermaInspect of light with wavelengths above 400 nm. Figure 6 shows images taken at a depth of approximately 90 μm inside the basal layer of a nevus. Spots which are significantly brighter than the surrounding region are clearly visible (Fig. 6 left). These spots also appear in the images taken at higher excitation wavelengths (Fig. 6 right), although, their brightness decreases significantly with increasing excitation wavelength. They can be assigned to the autofluorescence of melanin, which is abundant in the basal layer of nevi and has a one- photon-absorption maximum in the UV spectral range [4,7,27]. Figure 7 shows multiphoton images taken at a depth of 116 μm inside the skin. For wavelengths below λ = 800 nm two-photon autofluorescence of elastin and collagen structures, contribute to the signals in this layer. Above ~780 nm laser wavelength sufficient SHG light from the collagen structures reaches the PMT and becomes detectable and then accounts for the main contribution of the signal. The relative contributions of elastin and collagen, which can be used to determine the skin-aging index (SAAID) , can be distinguished by spectral separation. The autofluorescence of the elastin fibers themselves decreases with rising excitation wavelength.
Figure 8 summarizes the measured signal dependencies on the excitation wavelength of the fluorophores collagen, melanin, NAD(P)H, and keratin out of the different skin layers stratum corneum, stratum spinosum, stratum papillare, stratum basale, and the epidermal-dermal junction. The figure presents the mean intensities of the indicated regions of Fig. 3 - Fig. 6. In detail, the top-left graph shows the clear autofluorescence increase of keratin of the stratum corneum with decreasing excitation wavelength of the accessible spectral range which was already observed in Fig. 3. The apparent signal drop from λ = 740 nm to λ = 720 nm is an artifact of this image series which is possibly due to a small vertical movement of the skin region during the measurement of this series. Tiny movements < 3 µm in vertical direction between two images can particularly be notable in measurements of the stratum corneum because the tiny two-photon excitation volume can easily drop out of this thin top layer. The autofluorescence of NAD(P)H exhibits a trend similar to the keratin signal (Fig. 8, top right). As mentioned before the two-photon excitation of NAD(P)H for excitation wavelengths above λ = 800 nm has only very low efficiency. Therefore, the stratum spinosum can be best imaged with wavelengths between λ = 720 and λ = 780 nm. The Fig. 8 (bottom left) shows the signal intensity of melanin in a layer of the stratum basale. The maximum signal of melanin, which has a one-photon absorption maximum in the UV spectral range, is achieved within the measured wavelength range for an excitation wavelength of λ = 720 nm. For longer excitation wavelengths the signal significantly decreases. Nevertheless, it is more intense than the surrounding background even at λ = 880 nm (compare image series in Fig. 6). The multiphoton signal of collagen structures (Fig. 8, bottom right) strongly raises with increasing wavelengths and is within the measured wavelength range maximum at λ = 880 nm. The strong raise of the signal which is dominated by SHG light results from the wavelength dependencies of the generation and detection of the SHG light. For example, SHG light of wavelengths below 800 nm is hardly detected because of the low detection efficiency of the multiphoton tomograph for wavelengths in the UV spectral range. The data shown in Fig. 8 represent image gray values obtained with different mean powers and wavelength-dependent detection efficiencies. The signal intensities of the different fluorophores imaged at different depths cannot be directly derived because the excitation power had to be increased for imaging of the deeper skin layers due to the growing number of scattering events in tissue on the optical pathway. Nevertheless, Fig. 8 provides clear wavelength dependencies of two-photon autofluorescence and detected SHG signals of the major fluorophores of human skin measured in vivo.
3.3 In vivo measurements of collagen structures: distinction of SHG and autofluorescence collagen signals
The multiphoton images shown so far were obtained with the BG 39 color-glass filter inside the detection pathway, with no spectral separation of different signal contributions because of the broad transmission range of the filter. Figure 9 (a-c) shows multiphoton images of a 130 μm deep layer at the dermal-epidermal junction containing collagen structures with excitation wavelengths λ = 800 nm. The images were obtained by consecutively inserting the different filters BG 39, BP 395/11 and BP460/60, respectively, into the signal beam path. The signals can be assigned mainly to autofluorescence of elastin and SHG light from the collagen network [4,6]. By increasing the excitation wavelength to λ = 820 nm (Fig. 9 d), the mean intensity raises by ~10% with the BG 39 color-glass filter in place. This behavior results from a shift of the SHG wavelength from 400 nm to 410 nm, i.e. shifting to a wavelength that is more efficiently detected (compare Fig. 1). Figure 9 c) shows a multiphoton image at λ = 800 nm with 50 mW excitation power and the BP 460/60 which in this case blocks SHG light and transmits autofluorescence only. The image nicely shows detailed structures of elastin.
Thus, spectral filtering can be used to separate signals from elastin and collagen structures because for excitation wavelengths around λ = 800 nm the signal from collagen is exclusively SHG, whereas autofluorescence essentially originates from elastin . The separation of collagen SHG and elastin autofluorescence in dependence on laser wavelength is demonstrated in Fig. 10 . With the BG 39 color-glass filter both contributions are detected (squares). Using the BP 460/60 SHG light can be effectively suppressed for excitation wavelengths below λ = 830 nm leaving autofluorescence of elastin only (triangles). For larger excitation wavelengths, a 470-nm-longpass filter was used to suppress the SHG light leaving effectively again only the autofluorescence of elastin (circles) in the wavelength range from 800 nm - 920 nm. However, the detected SHG and autofluorescence contributions can also be distinguished by the time characteristics of the signal decay as described in the next section.
3.4 FLIM measurements
Figure 11 shows two FLIM images of the collagen and elastin structures already presented in Fig. 9, hence, obtained at a depth of 130 μm with an excitation wavelength of λ = 800 nm. To obtain the images the data of each pixel were previously fitted with a bi- (Fig. 11 left) and a mono-exponential decay curve (Fig. 11 right) with decay times τ1,τ2, and τ and amplitudes a1, a2 and a as fitting parameters, respectively. The values of τ2 and τ are false-color coded. The images were recorded with a BG 39 color-glass filter (Fig. 11, left) and the BP 460/60 (Fig. 11, right), respectively. The multiphoton excitation of this layer leads to both SHG and autofluorescence signals. While the BG 39 color-glass filter transmits both signal components, the bandpass filter blocks the 400-nm-SHG signal contribution of the collagen structures. Typical decay curves measured at marked pixels are shown in Fig. 12 . The black squares represent data measured with the BG 39 color-glass filter. Two signal contributions, a fast and a slower decaying component, are visible. The fast component originates from instantaneously generated SHG light with the temporal shape being spread out by the instrument response function. In the other case with the SHG light blocked the fitting reveals that the fast component accounts for about 97% of the total signal and the slower component for about 3% of the total signal. The remaining autofluorescence decays mainly mono-exponentially (Fig. 12, red circles). The fitting curves reveal decay times of τ2 = 2.1 ns (with SHG) and τ = 2.2 ns (SHG blocked).Figure 13 shows selected FLIM images of the stratum papillare with excitation wavelengths λ = 710 nm and λ = 800 nm, respectively. These images were obtained with the BG 39 color-glass filter. A fit with a two-exponential decay curve yielded amplitudes a1, a2 and decay times τ1, τ2. The decay times are presented by false color code in the figure. The decay curves clearly show a multi-exponential decay indicating the contributions of different fluorophores. In the images a filigree structure with several bright spots is visible which we assign mainly to melanin (an example is marked by a white circle). At an excitation wavelength of λ = 710 nm, no SHG light contributes to the signal because transmittance of the signal pathway is too low at 355 nm. Hence, the images of Fig. 13 (top row) result completely from the detection of autofluorescence light. With an excitation wavelength of λ = 800 nm the image considerably changes and a loop-like structure becomes visible. At λ = 800 nm the SHG signal due to collagen structures is detected and the account for the dominating contribution.
Several different fluorophores contribute to the signals at each pixel of the images of Fig. 13 which leads in dependence of the ratio of the fluorophores to a distribution of mean decay times for the whole image. The distribution of the decay times τ1 and τ2 is shown in Fig. 14 . Changing the excitation wavelengths changes the proportion of the signals of the different fluorophores and results in significant changes of the maximum values and width of the distributions of both τ1 and τ2. For λ = 710 nm the τ1 values are below 1.0 ns and the values of τ2 are broadly distributed around 2.7 ns ranging from 0.9 ns up to 5.0 ns. The broad distributions of decay times result from the mixture of contributions of different fluorophores including melanin and free and bound NAD(P)H as well as autofluorescence of collagen and elastin. For artificial melanin samples, values of τ1 = 0.04 ns and of τ2 = 1.2 ns were reported . Typical fluorescence lifetimes of free and bound NAD(P)H are 0.3 ns and 2.0 - 2.3 ns, respectively, as well as 0.3 ns / 2 ns for collagen and elastin . The red curves in Fig. 14 show the decay-time distribution with λ = 800 nm excitation. The decay time τ1 exhibits a sharp peak around 160 ps, which corresponds again to the instrument response function of the system and reflects the instantaneous SHG signal from collagen. The distribution of τ2 is significantly broader than that of τ1 with a maximum around 1.8 ns. The change in the distribution of τ1 and τ2 reflects the different fluorophore-excitation efficiencies and the change of the SHG wavelength. The FLIM images clearly confirm the results of the multiphoton images and illustrate how the change of the excitation wavelength leads to a change of the distributions of decay times due to the changing fractions of the different excited fluorophores.
4. Summary and conclusion
In summary, in vivo-multiphoton measurements of human skin with the multiphoton tomograph DermaInspect were presented. In particular, the dependence of the signal intensities on the excitation wavelengths was discussed for different skin layers. Clear trends of the autofluorescence signals arising from the endogenous fluorophores keratin, NAD(P)H, melanin, elastin and collagen as well as the corresponding trend of the detected SHG signal of collagen structures could be established in vivo. For keratin it was found that the autofluorescence signal in the stratum corneum clearly increases with decreasing excitation wavelength. A similar trend exhibits the autofluorescence signal of NAD(P)H, the major fluorophore in the stratum spinosum. The cells of the stratum spinosum can be imaged best with wavelengths between λ = 720 and λ = 780 nm. Melanin, in normal skin particularly present in the stratum basale, shows the strongest autofluorescence signal for an excitation wavelength of λ = 720 nm. The SHG signal could be clearly separated from the autofluorescence signal in spectral and in time-resolved measurements. The latter was demonstrated in decay curves derived from FLIM measurements, which were presented for measurement conditions with and without SHG contributions. The change of the signal proportions of the different fluorophores for different excitation wavelengths was further illustrated in FLIM images as well as in the resulting in different distributions of the decay times. It is anticipated that the results may help to further elucidate multiphoton imaging for in vivo clinical use.
We would like to thank L. Hatko and S. Evgenia for their vivid commitment and support for carrying out the measurements. This work was supported by the European Commission (Seventh Framework Programme, SKINSPECTION, HEALTH-F5-2008-201577).
References and links
3. P. T. So, C. Y. Dong, B. R. Masters, and K. M. Berland, “Two-photon excitation fluorescence microscopy,” Annu. Rev. Biomed. Eng. 2(1), 399–429 (2000). [CrossRef]
4. K. König, “Clinical multiphoton tomography,” J Biophotonics 1(1), 13–23 (2008). [CrossRef]
6. K. König and I. Riemann, “High-resolution multiphoton tomography of human skin with subcellular spatial resolution and picosecond time resolution,” J. Biomed. Opt. 8(3), 432–439 (2003). [CrossRef] [PubMed]
7. E. Dimitrow, I. Riemann, A. Ehlers, M. J. Koehler, J. Norgauer, P. Elsner, K. König, and M. Kaatz, “Spectral fluorescence lifetime detection and selective melanin imaging by multiphoton laser tomography for melanoma diagnosis,” Exp. Dermatol. 18(6), 509–515 (2009). [CrossRef] [PubMed]
8. E. Dimitrow, M. Ziemer, M. J. Koehler, J. Norgauer, K. König, P. Elsner, and M. Kaatz, “Sensitivity and specificity of multiphoton laser tomography for in vivo and ex vivo diagnosis of malignant melanoma,” J. Invest. Dermatol. 129(7), 1752–1758 (2009). [CrossRef] [PubMed]
9. M. J. Koehler, K. König, P. Elsner, R. Bückle, and M. Kaatz, “In vivo assessment of human skin aging by multiphoton laser scanning tomography,” Opt. Lett. 31(19), 2879–2881 (2006). [CrossRef] [PubMed]
10. M. J. Koehler, A. Preller, N. Kindler, P. Elsner, K. König, R. Bückle, and M. Kaatz, “Intrinsic, solar and sunbed-induced skin aging measured in vivo by multiphoton laser tomography and biophysical methods,” Skin Res. Technol. 15(3), 357–363 (2009). [CrossRef] [PubMed]
11. K. Schenke-Layland, I. Riemann, O. Damour, U. A. Stock, and K. König, “Two-photon microscopes and in vivo multiphoton tomographs--powerful diagnostic tools for tissue engineering and drug delivery,” Adv. Drug Deliv. Rev. 58(7), 878–896 (2006). [CrossRef] [PubMed]
12. F. Stracke, B. Weiss, C. M. Lehr, K. König, U. F. Schaefer, and M. Schneider, “Multiphoton microscopy for the investigation of dermal penetration of nanoparticle-borne drugs,” J. Invest. Dermatol. 126(10), 2224–2233 (2006). [CrossRef] [PubMed]
13. K. König, A. Ehlers, F. Stracke, and I. Riemann, “In vivo drug screening in human skin using femtosecond laser multiphoton tomography,” Skin Pharmacol. Physiol. 19(2), 78–88 (2006). [CrossRef] [PubMed]
14. J. Chen, S. Zhuo, T. Luo, X. Jiang, and J. Zhao, “Spectral characteristics of autofluorescence and second harmonic generation from ex vivo human skin induced by femtosecond laser and visible lasers,” Scanning 28(6), 319–326 (2006). [CrossRef] [PubMed]
15. J. A. Palero, H. S. de Bruijn, A. van der Ploeg van den Heuvel, H. J. Sterenborg, and H. C. Gerritsen, “Spectrally resolved multiphoton imaging of in vivo and excised mouse skin tissues,” Biophys. J. 93(3), 992–1007 (2007). [CrossRef] [PubMed]
16. C. Xu, W. Zipfel, J. B. Shear, R. M. Williams, and W. W. Webb, “Multiphoton fluorescence excitation: new spectral windows for biological nonlinear microscopy,” Proc. Natl. Acad. Sci. U.S.A. 93(20), 10763–10768 (1996). [CrossRef] [PubMed]
17. L. H. Laiho, S. Pelet, T. M. Hancewicz, P. D. Kaplan, and P. T. So, “Two-photon 3-D mapping of ex vivo human skin endogenous fluorescence species based on fluorescence emission spectra,” J. Biomed. Opt. 10(2), 024016 (2005). [CrossRef] [PubMed]
19. W. R. Zipfel, R. M. Williams, R. Christie, A. Y. Nikitin, B. T. Hyman, and W. W. Webb, “Live tissue intrinsic emission microscopy using multiphoton-excited native fluorescence and second harmonic generation,” Proc. Natl. Acad. Sci. U.S.A. 100(12), 7075–7080 (2003). [CrossRef] [PubMed]
20. P. T. C. S. Barry R. Masters, ed., Handbook of Biomedical Nonlinear Optical Microscopy (Oxford University Press, Inc., 2008).
21. R. M. Williams, W. R. Zipfel, and W. W. Webb, “Interpreting second-harmonic generation images of collagen I fibrils,” Biophys. J. 88(2), 1377–1386 (2005). [CrossRef]
22. M. S. Roberts, M. J. Roberts, T. A. Robertson, W. Sanchez, C. Thorling, Y. H. Zou, X. Zhao, W. Becker, and A. V. Zvyagin, “In vitro and in vivo imaging of xenobiotic transport in human skin and in the rat liver,” J. Biophoton. 1(6), 478–493 (2008). [CrossRef]
23. S. E. Cross, B. Innes, M. S. Roberts, T. Tsuzuki, T. A. Robertson, and P. McCormick, “Human skin penetration of sunscreen nanoparticles: in-vitro assessment of a novel micronized zinc oxide formulation,” Skin Pharmacol. Physiol. 20(3), 148–154 (2007). [CrossRef] [PubMed]
24. W. Becker, A. Bergmann, M. A. Hink, K. König, K. Benndorf, and C. Biskup, “Fluorescence lifetime imaging by time-correlated single-photon counting,” Microsc. Res. Tech. 63(1), 58–66 (2004). [CrossRef]
25. A. Pena, M. Strupler, T. Boulesteix, and M. Schanne-Klein, “Spectroscopic analysis of keratin endogenous signal for skin multiphoton microscopy,” Opt. Express 13(16), 6268–6274 (2005). [CrossRef] [PubMed]
26. A. Zoumi, A. Yeh, and B. J. Tromberg, “Imaging cells and extracellular matrix in vivo by using second-harmonic generation and two-photon excited fluorescence,” Proc. Natl. Acad. Sci. U.S.A. 99(17), 11014–11019 (2002). [CrossRef] [PubMed]
27. K. Teuchner, W. Freyer, D. Leupold, A. Volkmer, D. J. Birch, P. Altmeyer, M. Stücker, and K. Hoffmann, “Femtosecond two-photon excited fluorescence of melanin,” Photochem. Photobiol. 70(2), 146–151 (1999). [PubMed]
28. A. Zoumi, X. Lu, G. S. Kassab, and B. J. Tromberg, “Imaging coronary artery microstructure using second-harmonic and two-photon fluorescence microscopy,” Biophys. J. 87(4), 2778–2786 (2004). [CrossRef] [PubMed]
29. A. Ehlers, I. Riemann, M. Stark, and K. König, “Multiphoton fluorescence lifetime imaging of human hair,” Microsc. Res. Tech. 70(2), 154–161 (2007). [CrossRef]