JBO Letters

Sub‐40 fs, 1060‐nm Yb‐fiber laser enhances penetration depth in nonlinear optical microscopy of human skin

[+] Author Affiliations
Mihaela Balu, Jue Hou, Bruce J. Tromberg

University of California, Irvine, Beckman Laser Institute, Laser Microbeam and Medical Program, 1002 Health Sciences Road, Irvine, California 92612, United States

Ilyas Saytashev, Marcos Dantus

Michigan State University, Department of Chemistry, 578 South Shaw Lane, East Lansing, Michigan 48824, United States

J. Biomed. Opt. 20(12), 120501 (Dec 07, 2015). doi:10.1117/1.JBO.20.12.120501
History: Received June 29, 2015; Accepted October 30, 2015
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Open Access Open Access

Abstract.  Advancing the practical utility of nonlinear optical microscopy requires continued improvement in imaging depth and contrast. We evaluated second‐harmonic generation (SHG) and third‐harmonic generation images from ex vivo human skin and showed that a sub‐40 fs, 1060‐nm Yb‐fiber laser can enhance SHG penetration depth by up to 80% compared to a >100fs, 800 nm Ti:sapphire source. These results demonstrate the potential of fiber‐based laser systems to address a key performance limitation related to nonlinear optical microscopy (NLOM) technology while providing a low‐barrier‐to‐access alternative to Ti:sapphire sources that could help accelerate the movement of NLOM into clinical practice.

Figures in this Article

In vivo, label‐free nonlinear optical microscopy (NLOM) of human skin is under investigation for a broad range of clinical applications spanning from skin cancer detection and diagnosis14 to characterizing and understanding keratinocyte metabolism,5 skin aging,6,7 pigment biology,8,9 and cosmetic treatments.1012 NLOM signals are derived from several sources including cellular cofactors, melanin, and extracellular matrix proteins. Although exceptionally rich in both anatomic and functional contrast, NLOM has relatively limited penetration depth in turbid materials. This is due to the fact that multiple light scattering diminishes the instantaneous excitation intensity and nonlinear signal generation in the focused laser beam. As a result, there is considerable interest in exploring how light source performance can be optimized to improve imaging depth. Ti:sapphire lasers, commonly used in NLOM imaging, are generally able to access the superficial dermis of human skin to depths of 150 to 200μm. Penetration depth primarily depends on the material scattering length at the excitation wavelength, the efficiency of the nonlinear excitation process, the excitation average power, repetition rate, pulse width, and the detection geometry.13,14 Adjusting these parameters in order to improve the penetration depth has been explored in several studies using Ti:sapphire and optical parametric oscillator‐based femtosecond lasers as excitation light sources.1521 Sun et al. have shown that reduced light scattering using a Cr:Forsterite 1230 to 1250 nm source can increase penetration depth up to 300μm for harmonic generation imaging of human skin.22 Improvements in penetration depth can also be achieved when using shorter laser pulse widths.16 Depth resolved imaging studies require higher average laser powers and thermal damage to tissue becomes an issue of concern.23,24 Photothermal absorption of tissue is wavelength dependent, and so is the damage threshold. Heating following laser exposure at 800 nm is five times greater than at 1060 nm, and the damage threshold at 800 nm is three times lower than at 1060 nm.25 With the development of next‐generation fiber lasers, it is possible to imagine combining these technical features with more compact, portable, and inexpensive light sources that could facilitate clinical translation of NLOM technology.

Fiber‐based laser sources have been used for NLOM imaging of thin tissue cross‐sections,2628 mouse brain,29 and human skin tissue30 using fluorescence labeling. In this work, we evaluate the performance of a sub‐40 fs, 1060‐nm Yb‐fiber laser for label‐free NLOM imaging of human skin. The effect of excitation wavelength and pulse width on penetration depth in thick, turbid tissues is determined by comparing the fiber laser to an 800 nm Ti:sapphire laser source. We employ the depth‐dependent decay of second‐harmonic generation (SHG) signals as a standard metric for evaluating performance.

The excitation laser sources used were a Ti:sapphire oscillator (MIRA 900; Coherent Inc.; 220 fs, 76 MHz, 600 mW output power, tuning wavelength 720 to 980 nm) tuned to 800 nm for this study and a Yb‐fiber laser (BioPhotonic Solutions Inc., 1060 nm, sub‐40 fs, 39.2 MHz, 200 mW compressed output power). The prototype Yb‐fiber laser, with self‐similar pulse evolution,28 has an integrated adaptive phase‐amplitude pulse shaper (MIIPS‐HD, BioPhotonic Solutions Inc.) based on a 4f configuration with a two‐dimensional spatial light modulator. The purpose of the pulse shaper was to control high‐order phase distortions introduced by the high numerical‐aperture (NA) objective and other dispersive elements in the beam path. The 1060‐nm pulses were compressed to nearly transform limited duration using multiphoton intrapulse interference phase scan (MIIPS),31 and their full‐width half maximum duration was measured by interferometric autocorrelation using the microscope detection unit (BioPhotonic Solutions Inc.) at the focal plane. Each of the two excitation beams (800 and 1060 nm) was directed toward our home‐built laser‐scanning microscope and focused into the sample by an Olympus objective (XLPL25XWMP, 25×/1.05 NA water). The nonlinear signals from the sample were epi‐collected and directed toward two photomultiplier tubes (R3896, Hamamatsu) by a dichroic mirror (Semrock, Inc., 510 LP). The dichroic mirror was used to split the emission signal into two spectral channels defined by the emission filters: 440 SP; 375/110BP and 720 SP; 535/150BP (Semrock Inc.). We used discarded human skin tissue (fixed in formalin) to test the effect of sub‐40 fs, 1060‐nm excitation laser pulses on depth penetration in this sample. For each excitation wavelength (800 and 1060 nm), we acquired five stacks of images as optical sections of 430×430μm2 (512×512pixels) at different depths ranging from 0 to 200μm (2μm step). In the sample studied in this work, the main contrast mechanisms for 800‐nm excitation are based on two‐photon excited fluorescence (TPEF) signals from keratin, melanin, and elastin fibers and on SHG signals from collagen fibers. When using 1060 nm as excitation, the epidermis is visualized by third‐harmonic generation (THG) contrast derived from refractive index discontinuities at interfaces, while dermal contrast is derived from collagen fiber SHG.

Figure 1 shows merged images of human skin acquired at the same depth with 800‐ and 1060‐nm excitation. THG imaging of the keratinocyte structure in human skin epidermis using 1230 nm as excitation has been reported by Sun et al. in several studies.6,22 THG is not generated by elastin fibers in human dermis, although signals from elastic cartilage have been observed.32

Graphic Jump Location
Fig. 1
F1 :

Multimodal nonlinear optical microscopy (NLOM) images of human skin acquired with 800‐ and 1060‐nm excitation wavelengths at the same depth. (a) Epidermal‐dermal junction in human skin imaged by third‐harmonic generation (THG) (blue) and second‐harmonic generation (SHG) (red) using 1060 nm and by two‐photon excited fluorescence (TPEF) (green) using 800 nm as excitation wavelengths (z=35μm). TPEF signal originates from keratin in the epidermal keratinocytes and from elastin fibers (arrows) in the superficial papillary dermis, while THG signal highlights the keratinocytes only; SHG signal originates from collagen fibers. (b) Multimodal NLOM image corresponding to the inset in (a) representing contribution from three channels: (c) TPEF signal from keratinocytes and elastin fibers (arrows), (d) THG signal from keratinocytes, and (e) SHG signal from collagen fibers. Scale bar is 50μm.

Figure 2 shows representative images corresponding to one of the stacks acquired in the same location of the sample by using 800 and 1060 nm as excitation. The images in Figs. 2(a)2(c) and 2(f)2(h) represent en‐face (xy plane) images acquired at different depths. The cross‐sectional (xz plane) images shown in Figs. 2(d) and 2(e) were obtained from three‐dimensional (3‐D) image reconstruction of en‐face stacks using Amira (FEI Inc.).

Graphic Jump Location
Fig. 2
F2 :

Ex vivo imaging of human skin using 800 nm (Ti:sapphire laser) and 1060 nm (Yb‐fiber laser). (a–c) Horizontal sections (xy scans) at different depths corresponding to 800‐nm excitation wavelength. The optical sections show images of the epidermal cells through the TPEF signal (magenta, z=25μm); collagen fibers (green; SHG signal) and elastin fibers (magenta, TPEF signal) (z=100μm; 140μm). Vertical sections were obtained from three‐dimensional reconstruction for (d) 800‐nm and (e) 1060‐nm excitation wavelengths (40 mW for 800 nm and 20 mW for 1060 nm). Horizontal sections (xy scans) at different depths corresponding to 800‐ and 1060‐nm excitation wavelengths are shown in (a–c), (f–h), respectively. The optical sections show images of the epidermal cells through the THG signal (magenta, z=25μm) and collagen fibers (green; SHG signal) (z=100μm; 140μm). Scale bar is 50μm. The plot represents the SHG signal attenuation (logarithmic scale) with depth, for 800‐ and 1060‐nm excitation wavelengths.

To compare the penetration depth attained by each excitation wavelength, we adjusted the laser powers (40 mW for 800 nm and 20 mW for 1060 nm) such that the average intensity of the SHG signal corresponding to the sample surface (z=0) was similar for both wavelengths. The laser power and all the other experimental parameters were kept the same during the data acquisition. The SHG signals measured in the dermis of the skin sample are plotted versus depth in Fig. 2 on log scale. The signal calculated at each depth represents the average of the pixel intensities in the SHG images at that particular depth. The SHG intensity decay curve was normalized to the maximum intensity value for each wavelength.

The SHG intensity decays as a function of depth z according ISHGexp(Az), where A is the attenuation coefficient that includes the sample absorption and scattering properties at both the excitation and emission wavelengths. The inverse of A yields a 1/e attenuation length of 49μm for 800 nm and 88μm for 1060 nm, an increase of 80% for the Yb‐fiber laser source. Similar results were obtained for all five stacks acquired in the sample, which shows that 1060 nm, sub‐40 fs pulses can provide deeper penetration in highly scattering samples, such as skin.

In summary, these results demonstrate the potential of fiber‐based laser systems to be used as excitation light sources for NLOM imaging of highly turbid media. Despite their current lack of tunability, short‐pulse, >1μm wavelength fiber lasers can provide a low‐barrier‐to‐access alternative to conventional Ti:sapphire lasers. They are of particular interest in applications related to in vivo imaging of human skin as they can deliver up to 80% improvement in SHG imaging depth compared to conventionally used Ti:sapphire lasers. An additional benefit for in vivo human skin imaging is related to the THG contrast mechanism which, unlike TPEF, does not involve absorption and might allow for the use of higher excitation powers. With continued development of expanded wavelengths, powers, and pulse characteristics, these systems are expected to increase in use, particularly in skin studies where assessment of 3‐D morphology is important.

Acknowledgments

We would like to thank BioPhotonic Solutions Inc. for making their laser prototype available for these measurements and, in particular, Dr. Bingwei Xu for installing the laser system at UC Irvine. This research was supported partially by the National Institutes of Health (NIH) NIBIB Laser Microbeam and Medical Program (LAMMP, P41‐EB015890), Air Force Research Laboratory Agreement No. FA9550‐04‐1‐0101, and the Arnold and Mabel Beckman Foundation.

Balu  M.  et al., “Distinguishing between benign and malignant melanocytic nevi by in vivo multiphoton microscopy,” Cancer Res.. 74, (10 ), 2688 –2697 (2014). 0008‐5472 CrossRef
Balu  M.  et al., “In vivo multiphoton microscopy of basal cell carcinoma,” JAMA Dermatol.. 151, (10 ), 1068 –1074 (2015).CrossRef
Dimitrow  E.  et al., “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
Ulrich  M.  et al., “In vivo detection of basal cell carcinoma: comparison of a reflectance confocal microscope and a multiphoton tomography,” J. Biomed. Opt.. 18, (6 ), 061229  (2013). 1083‐3668 CrossRef
Balu  M.  et al., “In vivo multiphoton NADH fluorescence reveals depth‐dependent keratinocyte metabolism in human skin,” Biophys. J.. 104, (1 ), 258 –267 (2013). 0006‐3495 CrossRef
Liao  Y. H.  et al., “Quantitative analysis of intrinsic skin aging in dermal papillae by in vivo harmonic generation microscopy,” Biomed. Opt. Express. 5, (9 ), 3266 –3279 (2014). 2156‐7085 CrossRef
Koehler  M. J.  et al., “In vivo assessment of human skin aging by multiphoton laser scanning tomography,” Opt. Lett.. 31, (19 ), 2879 –2881 (2006). 0146‐9592 CrossRef
Krasieva  T. B.  et al., “Two‐photon excited fluorescence lifetime imaging and spectroscopy of melanins in vitro and in vivo,” J. Biomed. Opt.. 18, (3 ), 031107  (2013). 1083‐3668 CrossRef
Dancik  Y.  et al., “Use of multiphoton tomography and fluorescence lifetime imaging to investigate skin pigmentation in vivo,” J. Biomed. Opt.. 18, (2 ), 026022  (2013). 1083‐3668 CrossRef
Bazin  R.  et al., “Clinical study on the effects of a cosmetic product on dermal extracellular matrix components using a high‐resolution multiphoton tomography,” Skin Res. Technol.. 16, (3 ), 305 –310 (2010).CrossRef
Leite‐Silva  V. R.  et al., “The effect of formulation on the penetration of coated and uncoated zinc oxide nanoparticles into the viable epidermis of human skin in vivo,” Eur. J. Pharm. Biopharm.. 84, (2 ), 297 –308 (2013). 0939‐6411 CrossRef
Lademann  J.  et al., “In vivo methods for the analysis of the penetration of topically applied substances in and through the skin barrier,” Int. J. Cosmet. Sci.. 34, (6 ), 551 –559 (2012). 0142‐5463 CrossRef
Oheim  M  et al., “Two‐photon microscopy in brain tissue: parameters influencing the imaging depth,” J. Neurosci. Methods. 111, (1 ), 29 –37 (2001). 0165‐0270 CrossRef
Helmchen  F., and Denk  W., “Deep tissue two‐photon microscopy,” Nat. Methods. 2, (12 ), 932 –940 (2005). 1548‐7091 CrossRef
Theer  P., , Hasan  M. T., and Denk  W., “Two‐photon imaging to a depth of 1000 micron in living brains by use of a Ti:Al2O3 regenerative amplifier,” Opt. Lett.. 28, (12 ), 1022 –1024 (2003).CrossRef
Tang  S.  et al., “Effect of pulse duration on two‐photon excited fluorescence and second harmonic generation in nonlinear optical microscopy,” J. Biomed. Opt.. 11, (2 ), 020501  (2006). 1083‐3668 CrossRef
Balu  M.  et al., “Effect of excitation wavelength on penetration depth in nonlinear optical microscopy of turbid media,” J. Biomed. Opt.. 14, (1 ), 010508  (2009). 1083‐3668 CrossRef
Kobat  D.  et al., “Deep tissue multiphoton microscopy using longer wavelength excitation,” Opt. Express. 17, (16 ), 13354 –13364 (2009). 1094‐4087 CrossRef
Xi  P.  et al., “Two‐photon imaging using adaptive phase compensated ultrashort laser pulses,” J. Biomed. Opt.. 14, (1 ), 014002  (2009). 1083‐3668 CrossRef
Andresen  V.  et al., “Infrared multiphoton microscopy: subcellular‐resolved deep tissue imaging,” Curr. Opin. Biotechnol.. 20, (1 ), 54 –62 (2009). 0958‐1669 CrossRef
Kobat  D., , Horton  N. G., and Xu  C., “In vivo two‐photon microscopy to 1.6‐mm depth in mouse cortex,” J. Biomed. Opt.. 16, (10 ), 106014  (2011). 1083‐3668 CrossRef
Chen  S. Y., , Wu  H. Y., and Sun  C. K., “In vivo harmonic generation biopsy of human skin,” J. Biomed. Opt.. 14, (6 ), 060505  (2009). 1083‐3668 CrossRef
Masters  B. R.  et al., “Mitigating thermal mechanical damage potential during two‐photon dermal imaging,” J. Biomed. Opt.. 9, (6 ), 1265 –1270 (2004). 1083‐3668 CrossRef
Saytashev  I.  et al., “Pulse duration and energy dependence of photodamage and lethality induced by femtosecond near infrared laser pulses in Drosophila melanogaster,” J. Photochem. Photobiol. B. 115, , 42 –50 (2012). 1011‐1344 CrossRef
Bixler  J. N.  et al., “Assessment of tissue heating under tunable near‐infrared radiation,” J. Biomed. Opt.. 19, (7 ), 070501  (2014). 1083‐3668 CrossRef
Galli  R.  et al., “Non‐linear optical microscopy of kidney tumours,” J. Biophotonics. 7, (1–2 ), 23 –27 (2014).CrossRef
Galli  R.  et al., “Vibrational spectroscopic imaging and multiphoton microscopy of spinal cord injury,” Anal. Chem.. 84, (20 ), 8707 –8714 (2012).CrossRef
Nie  B.  et al., “Multimodal microscopy with sub‐30 fs Yb fiber laser oscillator,” Biomed. Opt. Express. 3, (7 ), 1750 –1756 (2012). 2156‐7085 CrossRef
Horton  N. G.  et al., “In vivo three‐photon microscopy of subcortical structures within an intact mouse brain,” Nat. Photonics. 7, (3 ), 205 –209 (2013). 1749‐4885 CrossRef
Tang  S.  et al., “Developing compact multiphoton systems using femtosecond fiber lasers,” J. Biomed. Opt.. 14, (3 ), 030508  (2009). 1083‐3668 CrossRef
Coello  Y.  et al., “Interference without an interferometer: a different approach to measuring, compressing, and shaping ultrashort laser pulses,” J. Opt. Soc. Am. B. 25, (6 ), A140 –A150 (2008). 0740‐3224 CrossRef
Yu  C. H.  et al., “In vivo and ex vivo imaging of intra‐tissue elastic fibers using third‐harmonic‐generation microscopy,” Opt. Express. 15, (18 ), 11167 –11177 (2007).CrossRef
© The Authors. Published by SPIE under a Creative Commons Attribution 3.0 Unported License. Distribution or reproduction of this work in whole or in part requires full attribution of the original publication, including its DOI.

Citation

Mihaela Balu ; Ilyas Saytashev ; Jue Hou ; Marcos Dantus and Bruce J. Tromberg
"Sub‐40 fs, 1060‐nm Yb‐fiber laser enhances penetration depth in nonlinear optical microscopy of human skin", J. Biomed. Opt. 20(12), 120501 (Dec 07, 2015). ; http://dx.doi.org/10.1117/1.JBO.20.12.120501


Figures

Graphic Jump Location
Fig. 1
F1 :

Multimodal nonlinear optical microscopy (NLOM) images of human skin acquired with 800‐ and 1060‐nm excitation wavelengths at the same depth. (a) Epidermal‐dermal junction in human skin imaged by third‐harmonic generation (THG) (blue) and second‐harmonic generation (SHG) (red) using 1060 nm and by two‐photon excited fluorescence (TPEF) (green) using 800 nm as excitation wavelengths (z=35μm). TPEF signal originates from keratin in the epidermal keratinocytes and from elastin fibers (arrows) in the superficial papillary dermis, while THG signal highlights the keratinocytes only; SHG signal originates from collagen fibers. (b) Multimodal NLOM image corresponding to the inset in (a) representing contribution from three channels: (c) TPEF signal from keratinocytes and elastin fibers (arrows), (d) THG signal from keratinocytes, and (e) SHG signal from collagen fibers. Scale bar is 50μm.

Graphic Jump Location
Fig. 2
F2 :

Ex vivo imaging of human skin using 800 nm (Ti:sapphire laser) and 1060 nm (Yb‐fiber laser). (a–c) Horizontal sections (xy scans) at different depths corresponding to 800‐nm excitation wavelength. The optical sections show images of the epidermal cells through the TPEF signal (magenta, z=25μm); collagen fibers (green; SHG signal) and elastin fibers (magenta, TPEF signal) (z=100μm; 140μm). Vertical sections were obtained from three‐dimensional reconstruction for (d) 800‐nm and (e) 1060‐nm excitation wavelengths (40 mW for 800 nm and 20 mW for 1060 nm). Horizontal sections (xy scans) at different depths corresponding to 800‐ and 1060‐nm excitation wavelengths are shown in (a–c), (f–h), respectively. The optical sections show images of the epidermal cells through the THG signal (magenta, z=25μm) and collagen fibers (green; SHG signal) (z=100μm; 140μm). Scale bar is 50μm. The plot represents the SHG signal attenuation (logarithmic scale) with depth, for 800‐ and 1060‐nm excitation wavelengths.

Tables

References

Balu  M.  et al., “Distinguishing between benign and malignant melanocytic nevi by in vivo multiphoton microscopy,” Cancer Res.. 74, (10 ), 2688 –2697 (2014). 0008‐5472 CrossRef
Balu  M.  et al., “In vivo multiphoton microscopy of basal cell carcinoma,” JAMA Dermatol.. 151, (10 ), 1068 –1074 (2015).CrossRef
Dimitrow  E.  et al., “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
Ulrich  M.  et al., “In vivo detection of basal cell carcinoma: comparison of a reflectance confocal microscope and a multiphoton tomography,” J. Biomed. Opt.. 18, (6 ), 061229  (2013). 1083‐3668 CrossRef
Balu  M.  et al., “In vivo multiphoton NADH fluorescence reveals depth‐dependent keratinocyte metabolism in human skin,” Biophys. J.. 104, (1 ), 258 –267 (2013). 0006‐3495 CrossRef
Liao  Y. H.  et al., “Quantitative analysis of intrinsic skin aging in dermal papillae by in vivo harmonic generation microscopy,” Biomed. Opt. Express. 5, (9 ), 3266 –3279 (2014). 2156‐7085 CrossRef
Koehler  M. J.  et al., “In vivo assessment of human skin aging by multiphoton laser scanning tomography,” Opt. Lett.. 31, (19 ), 2879 –2881 (2006). 0146‐9592 CrossRef
Krasieva  T. B.  et al., “Two‐photon excited fluorescence lifetime imaging and spectroscopy of melanins in vitro and in vivo,” J. Biomed. Opt.. 18, (3 ), 031107  (2013). 1083‐3668 CrossRef
Dancik  Y.  et al., “Use of multiphoton tomography and fluorescence lifetime imaging to investigate skin pigmentation in vivo,” J. Biomed. Opt.. 18, (2 ), 026022  (2013). 1083‐3668 CrossRef
Bazin  R.  et al., “Clinical study on the effects of a cosmetic product on dermal extracellular matrix components using a high‐resolution multiphoton tomography,” Skin Res. Technol.. 16, (3 ), 305 –310 (2010).CrossRef
Leite‐Silva  V. R.  et al., “The effect of formulation on the penetration of coated and uncoated zinc oxide nanoparticles into the viable epidermis of human skin in vivo,” Eur. J. Pharm. Biopharm.. 84, (2 ), 297 –308 (2013). 0939‐6411 CrossRef
Lademann  J.  et al., “In vivo methods for the analysis of the penetration of topically applied substances in and through the skin barrier,” Int. J. Cosmet. Sci.. 34, (6 ), 551 –559 (2012). 0142‐5463 CrossRef
Oheim  M  et al., “Two‐photon microscopy in brain tissue: parameters influencing the imaging depth,” J. Neurosci. Methods. 111, (1 ), 29 –37 (2001). 0165‐0270 CrossRef
Helmchen  F., and Denk  W., “Deep tissue two‐photon microscopy,” Nat. Methods. 2, (12 ), 932 –940 (2005). 1548‐7091 CrossRef
Theer  P., , Hasan  M. T., and Denk  W., “Two‐photon imaging to a depth of 1000 micron in living brains by use of a Ti:Al2O3 regenerative amplifier,” Opt. Lett.. 28, (12 ), 1022 –1024 (2003).CrossRef
Tang  S.  et al., “Effect of pulse duration on two‐photon excited fluorescence and second harmonic generation in nonlinear optical microscopy,” J. Biomed. Opt.. 11, (2 ), 020501  (2006). 1083‐3668 CrossRef
Balu  M.  et al., “Effect of excitation wavelength on penetration depth in nonlinear optical microscopy of turbid media,” J. Biomed. Opt.. 14, (1 ), 010508  (2009). 1083‐3668 CrossRef
Kobat  D.  et al., “Deep tissue multiphoton microscopy using longer wavelength excitation,” Opt. Express. 17, (16 ), 13354 –13364 (2009). 1094‐4087 CrossRef
Xi  P.  et al., “Two‐photon imaging using adaptive phase compensated ultrashort laser pulses,” J. Biomed. Opt.. 14, (1 ), 014002  (2009). 1083‐3668 CrossRef
Andresen  V.  et al., “Infrared multiphoton microscopy: subcellular‐resolved deep tissue imaging,” Curr. Opin. Biotechnol.. 20, (1 ), 54 –62 (2009). 0958‐1669 CrossRef
Kobat  D., , Horton  N. G., and Xu  C., “In vivo two‐photon microscopy to 1.6‐mm depth in mouse cortex,” J. Biomed. Opt.. 16, (10 ), 106014  (2011). 1083‐3668 CrossRef
Chen  S. Y., , Wu  H. Y., and Sun  C. K., “In vivo harmonic generation biopsy of human skin,” J. Biomed. Opt.. 14, (6 ), 060505  (2009). 1083‐3668 CrossRef
Masters  B. R.  et al., “Mitigating thermal mechanical damage potential during two‐photon dermal imaging,” J. Biomed. Opt.. 9, (6 ), 1265 –1270 (2004). 1083‐3668 CrossRef
Saytashev  I.  et al., “Pulse duration and energy dependence of photodamage and lethality induced by femtosecond near infrared laser pulses in Drosophila melanogaster,” J. Photochem. Photobiol. B. 115, , 42 –50 (2012). 1011‐1344 CrossRef
Bixler  J. N.  et al., “Assessment of tissue heating under tunable near‐infrared radiation,” J. Biomed. Opt.. 19, (7 ), 070501  (2014). 1083‐3668 CrossRef
Galli  R.  et al., “Non‐linear optical microscopy of kidney tumours,” J. Biophotonics. 7, (1–2 ), 23 –27 (2014).CrossRef
Galli  R.  et al., “Vibrational spectroscopic imaging and multiphoton microscopy of spinal cord injury,” Anal. Chem.. 84, (20 ), 8707 –8714 (2012).CrossRef
Nie  B.  et al., “Multimodal microscopy with sub‐30 fs Yb fiber laser oscillator,” Biomed. Opt. Express. 3, (7 ), 1750 –1756 (2012). 2156‐7085 CrossRef
Horton  N. G.  et al., “In vivo three‐photon microscopy of subcortical structures within an intact mouse brain,” Nat. Photonics. 7, (3 ), 205 –209 (2013). 1749‐4885 CrossRef
Tang  S.  et al., “Developing compact multiphoton systems using femtosecond fiber lasers,” J. Biomed. Opt.. 14, (3 ), 030508  (2009). 1083‐3668 CrossRef
Coello  Y.  et al., “Interference without an interferometer: a different approach to measuring, compressing, and shaping ultrashort laser pulses,” J. Opt. Soc. Am. B. 25, (6 ), A140 –A150 (2008). 0740‐3224 CrossRef
Yu  C. H.  et al., “In vivo and ex vivo imaging of intra‐tissue elastic fibers using third‐harmonic‐generation microscopy,” Opt. Express. 15, (18 ), 11167 –11177 (2007).CrossRef

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