Journal of BIOPHOTONICSbiophotonics.illinois.edu/pubs/biophotonics_current/JBP... · 2018-04-24 · FULL ARTICLE Integrated multimodal optical microscopy for structural and functional

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Integrated multimodal optical microscopy for structuraland functional imaging of engineered and natural skin

Youbo Zhao1, Benedikt W. Graf 1; 2, Eric J. Chaney1, Ziad Mahmassani3, Eleni Antoniadou4,Ross DeVolder5, Hyunjoon Kong5, Marni D. Boppart3, and Stephen A. Boppart*; 1; 2; 4; 6

1 Biophotonics Imaging Laboratory, Beckman Institute for Advanced Science and Technology, University of Illinois at Urbana-Champaign,Urbana, IL 61801, USA

2 Department of Electrical and Computer Engineering, University of Illinois at Urbana-Champaign, Urbana, IL 61801, USA3 Department of Kinesiology and Community Health, University of Illinois at Urbana-Champaign, Urbana, IL 61801, USA4 Department of Bioengineering, University of Illinois at Urbana-Champaign, Urbana, IL 61801, USA5 Department of Chemical and Biomolecular Engineering, University of Illinois at Urbana-Champaign, Urbana, IL 61801, USA6 Department of Internal Medicine, University of Illinois at Urbana-Champaign, Urbana, IL 61801, USA

Received 4 January 2012, revised 7 February 2012, accepted 7 February 2012Published online 28 February 2012

Key words: skin, tissue engineering, microscopy, functional imaging, multiphoton fluorescence microscopy, fluorescencelifetime imaging microscopy, optical coherence tomography

# 2012 by WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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An integrated multimodal optical microscope is demon-strated for high-resolution, structural and functional im-aging of engineered and natural skin. This microscopeincorporates multiple imaging modalities including opticalcoherence (OCM), multi-photon (MPM), and fluores-cence lifetime imaging microscopy (FLIM), enabling si-multaneous visualization of multiple contrast sourcesand mechanisms from cells and tissues. Spatially co-re-gistered OCM/MPM/FLIM images of multi-layered skintissues are obtained, which are formed based on comple-mentary information provided by different modalities,i.e., scattering information from OCM, molecular infor-mation from MPM, and functional cellular metabolismstates from FLIM. Cellular structures in both the dermisand epidermis, especially different morphological andphysiological states of keratinocytes from different epi-dermal layers, are revealed by mutually-validatingimages. In vivo imaging of human skin is also investi-gated, which demonstrates the potential of multimodalmicroscopy for in vivo investigation during engineeredskin engraftment. This integrated imaging technique and

microscope show the potential for investigating cellulardynamics in developing engineered skin and following invivo grafting, which will help refine the control and cul-turing conditions necessary to obtain more robust andphysiologically-relevant engineered skin substitutes.

Multimodal microscopy images of a microporous 3D hy-drogel scaffold seeded with 3T3 fibroblasts. Representa-tive spatially co-registered images were generated basedon different methodologies including optical coherence(OCM), multiphoton (MPM), and fluorescence lifetimeimaging (FLIM) microscopy.

* Corresponding author: e-mail: [email protected], Phone: +1 217 244 7479, Fax: +1 217 333 5833

J. Biophotonics 5, No. 5–6, 437–448 (2012) / DOI 10.1002/jbio.201200003

1. Introduction

Engineered skin equivalents have become importantbiomaterials for both basic research and clinicalpractice [1, 2]. Research-based applications ofsynthesized skin tissues, as an alternative to animaltesting, have included tests of efficacy of pharmaceu-tical and cosmetic products [3], studies of toxicitiesof chemical agents [4], understanding of normal skinfunction (e.g. UV protection [5], wound healing) anddermatological disease pathogenesis [6]. Clinically,engraftment of engineered skin has emerged as apromising option for the treatment of skin loss, espe-cially in the case of large-area injury or loss whenautologous transplant sources are not sufficient [7,8]. For all these applications, skin equivalents thathave well-developed three-dimensional (3D) struc-tures and mimic as many of the physiological andmechanical functions of native skin as possible, arealways preferred.

Tissue engineering of skin is a delicate multistepprocedure involving complex cell dynamics which in-terplay with culturing conditions and culturing mate-rials [9]. The use of 3D topographic scaffold struc-tures at the dermal-epidermal junction and theapplication of mechanical stimuli [10, 11] applied todeveloping tissues have been found to be importantfactors influencing the viability and functionality ofengineered skin tissues. It has been shown that acapable multi-functional imaging instrument wouldgreatly facilitate not only the investigation of devel-oping engineered skin tissues, but also the clinicalevaluation of successful in vivo engraftment, andthus help to evaluate and optimize the necessary cul-turing conditions to improve the applicability andoutcome of engineered skin tissues in the treatmentof skin injury, loss, and disease [12, 13].

Skin, including engineered skin constructs, is aspatially heterogeneous tissue with large variationsin cellular composition and organization, both instates of health and disease. An advantageous im-aging tool for skin is one that has sufficient resolutionand contrast to directly observe both the structuraland physiological or functional properties at differ-ent skin depths. Histology, the “gold standard”, hasbeen well-suited for the purpose of assessing cellmorphology and organization, and other architectu-ral aspects such as epidermal and dermal thickness,and orientation of collagen. However, histology pro-cedures are destructive, and always the endpoint forthe tissue under examination, making histology inap-propriate for the evaluation of the real-time cellulardynamics and physiological functions of the skin. Ul-trasound is non-invasive, but poorly suited for im-aging skin because of its low resolution, typically onthe order of 100 mm. High-frequency ultrasound ima-ging can achieve higher imaging resolution, such asless than 10 mm at 100 MHz frequency [14, 15], but is

not commonly used for skin imaging, and requirescontact and the use of an impedance-matching gel.

Optical coherence tomography (OCT) [16, 17] isan emerging biomedical imaging technology that hasunique features, such as non-contact, non-invasivedetection, higher resolution than ultrasound, betterimaging depth than conventional confocal micro-scopy, and has been widely employed in the field oftissue engineering [11, 18, 19]. The limitations ofOCT, however, include the relatively low resolutioncompared to histology, and deficient endogenouscontrast when used without contrast agents. Non-linear optical microscopy technologies, such as two-photon excitation fluorescence (TPEF) [20], secondharmonic generation (SHG) [21], and coherent anti-Stokes Raman scattering (CARS) [22], provide sub-cellular resolution and unique contrast capabilities.These advanced microscopy techniques have drawnconsiderable attention from dermatologists [13, 23].However, these microscopy techniques are plaguedby shallow imaging depths and somewhat specific orlimited contrast generation.

Given the different strengths and weaknesses as-sociated with these different technologies or modal-ities, a great deal of research effort has been investedin the development of integrated microscopes thatcombine multiple modalities and enable co-regis-tered images to leverage the advantages from multi-ple structural and functional contrast mechanisms.Multimodal microscopes such as an integrated multi-photon (MPM, including TPEF and SHG)/OCT [24],CARS/OCT [25], CARS/MPM [26], reflective confo-cal/TPEF/SHG [27, 28], fluorescence lifetime im-aging microscopy (FLIM)/SHG/TPEF [29, 30] amongothers, have been demonstrated successfully for si-multaneously visualizing multiple properties of thesample or specimen. In particular, the combinationof optical coherence microscopy (OCM, the high-re-solution variant of OCT [31]) and MPM has drawnconsiderable attention because this combination canprovide co-registered structural and molecular infor-mation of the sample and possibly be based on a sin-gle laser source [32, 33]. Other than this combination,an integrated OCT and FLIM system has been de-monstrated for simultaneous characterization of mor-phological and biochemical properties of tissue [34].

Fluorescence lifetime imaging microscopy is anestablished imaging methodology which is used toprobe molecular level variations associated with bio-logical activities, based on the changes in fluores-cence lifetime of specific fluorophores [30, 35, 36]. Ithas been shown to be capable of monitoring protein-protein interactions [37], detecting environmentalchanges in biological systems (such as pH value, ionconcentration, etc.), detecting early changes asso-ciated with certain cancers [38], and differentiatingapoptosis and necrosis pathways [39, 40]. In particu-lar, FLIM is a non-invasive method for monitoring

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metabolic states of living cells, based on the lifetimevariation of autofluorescence from reduced nicotin-amide adenine dinucleotide (NADH) [39, 41]. Acombination of FLIM with OCT has the potentialclinical application for early detection of pathologi-cal changes in biological tissues. Unfortunately, be-cause the two modalities have different workingprinciples, the information provided by the two mod-alities is not completely spatially co-registered, i.e.,3D volumetric OCT morphology versus two-dimen-sional superficial maps of autofluorescence lifetime.

Here we report an integrated imaging platformwhich incorporates MPM, OCM, and FLIM, conti-nuing our previous efforts for developing multimodalmicroscopy techniques [32, 42]. The combination ofFLIM not only adds a new contrast capability to ourpreviously developed OCM/MPM microscope, butmore importantly enables one to monitor the meta-bolic activities of living tissue. This new multimodal-enabled integrated microscope is used to character-ize multiple properties of tissue, based on an increas-ingly comprehensive set of information, i.e., scatter-ing information from OCM, molecular informationfrom MPM, and cellular metabolism from FLIM.

In this paper, real-time imaging of engineeredand natural skin is demonstrated with this multimo-dal microscope. For well-developed bi-layer (dermisand epidermis) engineered skin tissues, detailed mor-phology of cells and structures from different layerswere readily visualized with OCM and MPM, includ-ing keratinocytes in the epidermis, and fibroblastsand collagen in the dermis. Changes in morphologyand the metabolic state of keratinocytes from differ-ent epidermal layers, which represents the complex,multi-step cornification process of keratinocytes,were observed with OCM and FLIM. The keratini-zation process was also manifested by time-lapseimaging with OCM and MPM. Complementary con-trast capabilities were investigated by imaging fibro-blasts seeded in 3D structured hydrogel scaffolds. Invivo imaging of human skin was also investigated,which explores the capability of this system fortracking the skin biology following engraftment ofengineered skin tissue. Collectively, the use of a mul-timodal microscopy techniques in a single integratedimaging platform has the potential to improve ourunderstanding and diagnostic capabilities in derma-tology, and the development of engineered skin fortissue replacement.

2. Experiments and methods

In brief, the integrated OCM/MPM/FLIM imagingsystem is based on a single laser source, and ashared microscope frame, but with separated excita-tion beams and detection channels for different

modalities. In our previous work on the develop-ment of an integrated OCM/MPM microscope, sev-eral unique strategies were implemented for seam-less integration of different modalities. Thesestrategies include a single laser source based on adual-band arrangement, independently controlledexcitation beams, and computational algorithms forcorrection of coherence curvature, for dispersionbalance and correction, and for fast signal and imageprocessing. With these techniques, we are able to op-timize the performance of multiple imaging modal-ities simultaneously without compromising the ad-vantages of the individual modes. Most of themethodologies presented here are inherited fromour previous work, and detailed information can befound elsewhere [32, 43, 44].

The schematic of the integrated OCM/MPM/FLIM microscope is shown in Figure 1. For its opti-cal source, a high-power, widely-tunable titanium:-sapphire laser (Mai-Tai HP, Spectra-Physics) is used,

Figure 1 (online color at: www.biophotonics-journal.org)Schematic of the integrated optical microscope. Abbrevia-tions: BS, beam splitter; C, collimator; CCD, charge-coupled device line-scan camera; CL, coupling lens; DC,dichroic mirror; DG, diffraction grating; F, filter; GS, gal-vanometer scanner; HWP, half-wave plate; L, lens; O, ob-jective; M, high reflection mirror; PCF, photonic crystal fi-ber; PBS, polarizing beam splitter; PH, pinhole; PMT,photo-multiplier tube; RM, reference mirror; S, sample,TL, telescope lens.

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which outputs 100 fs pulses at an 80 MHz repetitionrate with a bandwidth of 10 nm at a tunable centerwavelength within the range of 730–1000 nm. Thelinearly-polarized output (with maximum averagepower of 3 W) from this laser is divided by a 90/10beam-splitter into two portions. The 10% portion,with narrowband but wide wavelength tuning range,is used for excitation of MPM and FLIM. The higherpower beam is coupled into a photonic crystal fiber(LMA-PM-5, NKT Photonics) where through self-phase modulation, the spectrum is broadened to100–150 nm FWHM, depending on the center wave-length of the laser. The light from the fiber, whichserves as the OCM source, is collimated with an ob-jective and then directed to the OCM interferom-eter. The polarization-maintaining fiber ensures thatthe OCM beam remains linearly polarized. TheOCM sample arm and the MPM/FLIM beams arecombined by a polarizing beam splitter, with polari-zation of the MPM/FLIM being controlled by a halfwave plate. The power of the MPM and OCMsources is independently controlled by a set of neu-tral density filters.

The collinearly-propagating OCM sample armand MPM beams are expanded by a telescope andfocused by a microscope objective (XLUMP20X,0.95 NA, water immersion, Olympus) into the sam-ple. While the back aperture of the objective is fullyfilled by the MPM beam to obtain an optimal resolu-tion, the beam diameter of the OCM sample armbeam is controlled by using a different collimator orputting a telescope right after the collimator. Bydoing this, we are able to control the effective NA ofthe OCM beam to obtain an optimal balance be-tween resolution and depth-of-field. For all theOCM images in this paper, if not specifically men-tioned, we used effective NA of 0.65. The sample ispositioned on a motorized stage which can translatethe sample in three directions. A pair of galvan-ometers (Micromax 671, Cambridge Technology) po-sitioned before the telescope is used to scan thebeam across the sample. The spectral interferencepattern of the reference and sample arm beams isdetected for OCM acquisition by a spectrometerwhich is based on a diffraction grating and CCD linecamera (Piranha2 2k, Dalsa). The frame rate for theline camera depends on the speed and mode of im-age acquisition (galvanometer), and can be up to32 kHz. OCM images are generated after severalprocessing steps including compensating for unba-lanced dispersion in the sample arm, compensatingfor non-uniform distribution of the spectrum on theCCD due to nonlinearity of the diffraction grating,and correction for the coherence curvature due tothe mismatch between the coherence gate and theconfocal gate across the field-of-view [32].

For MPM and FLIM, the epi-collected signals arediverted by a long-pass dichroic mirror and band-

pass filtered. The dichroic mirror can be flipped toopposite directions to steer the fluorescence signalsfor either MPM or FLIM. For MPM, the signal iseither divided by the dichroic mirrors or filtered intodifferent channels for parallel detection of variousfluorescence or SHG signals. The fluorescence sig-nals are detected by photomultiplier tubes (H7421,Hamamatsu). For FLIM, the fluorescence signals arecollected by a multimode fiber and then coupled intoa grating-based polychromator and detected by a 16-channel PMT. The output pulses of the PMT chan-nels are digitized and analyzed by a time-correlatedsingle-photon counting board (SPC 150, Becker &Hickle). FLIM analyses are performed by commer-cial software (SPCimage, Becker & Hickle). All theFLIM images in this paper, if not specifically men-tioned, were obtained by first integrating all 16 chan-nels and then fitting the appropriate lifetime decay.

Full thickness bi-layer engineered skin tissues(EpiDerm FT400, MatTek Corp.) were used to de-monstrate the performance of this integrated micro-scope. The engineered skin consists of normal, hu-man-derived epidermal keratinocytes, and normal,human-derived dermal fibroblasts, which form amulti-layered, highly differentiated model of the hu-man dermis and epidermis. The epidermis consists ofmostly keratinocytes of different physiological statesand the dermal compartment is composed of a col-lagen matrix containing viable human fibroblasts.The complicated multi-layered structure and physio-logically-rich cellular state make this engineered skinan excellent sample to test the imaging and diagnos-tic capabilities of our system. The tissues were cul-tured in an incubator with 5% CO2 at 37 �C, andnourished daily with 2.5 ml Dulbecco’s ModifiedEagle’s Medium (pH 7.4). The tissues were main-tained in single well tissue culture plate inserts (Cost-ar SnapwellTM, Corning) and at an air-liquid inter-face. During imaging, the engineered tissues werekept in a Petri dish with fresh culturing media atroom temperature. The Petri dish was slightly modi-fied with embedded stainless steel pins to hold the tis-sues and limit movement during imaging. The Petridish was also covered by a specially designed insertwith an optical window to help maintain the sterileculture environment of the tissue during imaging.After each imaging session, which required less thantwo hours, the tissues were returned back to their ori-ginal culturing conditions in a commercial incubator.

Fibroblasts seeded in microporous hydrogel scaf-folds were also used in these experiments, and wereprepared as follows. Alginate (LF20/40, FMC BioPo-lymer, Rockland, ME), weight-average moleculeweight (MW) �250,000 g mol�1, was dissolved in a0.1 M MES (2-(N-morpholino)ethanesulfonic acid)buffer (pH 6.4, Sigma-Aldrich, St. Louis, MO). The2% (w/w) sodium alginate solution was filtered and1-hydroxybenzotriazole (HOBt, Fluka, St. Louis,


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MO), 1-ethyl-3-(3-dimethylaminopropyl) carbodi-imide (EDC, Thermo Scientific, Hanover Park, IL),and adipic acid dihydrazide (AAD), Sigma-Aldrich)were dissolved in the alginate solution, and the pHwas adjusted to 6.0. The molar ratio of EDC : HOBt: AAD was 1.0 : 0.5 : 0.2. Subsequently, gel sheetswere made on glass slides and hydrogel disks withdiameters of 5 mm were punched out and incubatedin phosphate-buffered saline, at pH 7.4 (PBS) and at37 �C for 6 hr before characterization. The watercontent of the initial cross-linked hydrogel was97 wt%, which was determined by measuring the hy-drogel weight before and after swelling at 37 �C inmillipore water (Nanopure DiamondTM, Barnstead).The disks were then frozen at �20 �C for 6 hr andlyophilized at �60 �C. Finally, the gel discs were irra-diated with UV light (lmax ¼ 254 nm, Jelight Co.Model 20, Irvine, CA) for 2 hr for sterilization pur-poses. NIH 3T3 fibroblasts labeled with CellLightReagents (Life Technologies) were suspended in thepre-gel solution at a density of 1 � 106 cells ml�1.Cell viability was checked with MTT assay analysisat different time points within a few days of cellseeding.

Dorsal hand skin of a human volunteer served asthe natural skin specimen for in vivo imaging experi-ments. An imaging mount, which consisted of a thinglass coverslip mounted to a rigid holder, was de-signed to keep the skin stabilized during image ac-quisition. A small amount of glycerol was applied tothe skin prior to imaging to serve as an index match-ing agent. All imaging was done in accordance witha protocol approved by the Institutional ReviewBoard at the University of Illinois at Urbana-Cham-paign.

3. Results

Depth-resolved OCM/MPM images of a full thick-ness engineered skin sample are shown in Figure 2.All images, composed of 256 � 256 pixels, were ob-tained at a rate of about 3 seconds per frame. Allthe OCM and MPM images in this paper, if not spe-cifically mentioned, were generated with the sameparameters, and without any image averaging orprocessing. All the images at different depths (fromall three modalities) were obtained by axially trans-lating the samples via the three-axis translation stagewith the incidence laser power left unchanged. TheOCM images were generated based on the back-scattered light from the tissue, and the OCM depthresolution was determined by the coherence gatebased on the same principle as in OCT. In our case,the depth resolution was estimated to be 1.5 mm,which was determined by the 150 nm bandwidth ofthe supercontinuum-generated OCM beam from the

photonic crystal fiber. The signals in the MPMimages are primarily two-photon excited autofluo-rescence of NADH from cells and second harmonicgeneration from collagen. The axial and transverseresolutions for MPM were experimentally deter-mined by imaging fluorescent beads and were meas-ured to be 0.8 mm and 0.5 mm, respectively. Whileacquiring the MPM images, a short-pass filter wasused, and the autofluorescence and SHG signalswere not separated. The input laser energies forOCM and MPM were 5 and 22 mW, respectively.The center wavelengths used for OCM and MPMwere 800 nm and 760 nm, respectively. A histologyimage is shown in Figure 2(b) for comparison. Theco-registration of images from the two differentmodalities is shown by the merged image in Fig-ure 2(c).

Figure 2 (online color at: www.biophotonics-journal.org)OCM and MPM images from a bi-layer engineered skinsample. (a) OCM (left column) and MPM (right column)images from different layers of the engineered skin tissue.(b) Hematoxylin and eosin stained histology image of thetissue. (c) Pseudo-color images of OCM (top), MPM (mid-dle) and merged OCM and MPM (bottom). Abbrevia-tions: SC, stratum corneum; SG, stratum granulosum; SS,stratum spinosum; SB, stratum basale; BM, basal mem-brane; D, dermis. Scale bars represent 20 mm.




As shown by the OCM and MPM images in Fig-ure 2, different cellular architecture at differentdepths is visualized, including different morphologiesof keratinocytes from different layers, i.e., stratumcorneum (SC, stratified keratinocytes), stratum gran-ulosum (SG), stratum spinosum (SS), stratum basale(SB), as well as collagen and fibroblasts in the der-mis. Based on the different physics for contrast andimage formation, the two modalities provide differ-ent but complementary information of the sample.In particular, the morphological changes of the kera-tinocytes progressing from the inner layers to theouter layers reflect the evolution of the physiologicalstate at different points along the complicated pro-cess of keratinization. It is also worthwhile to notethat the SHG signals from the collagen are also vi-tally important information about the health and in-tegrity of skin [45], and have been used to investi-gate dermal diseases in vivo that are associated withchanges in collagen and elastin such as sclerodermaor graft versus host disease [13].

As a potentially non-invasive imaging method,another advantage of this integrated microscopy isfor the real-time monitoring of the dynamics or de-velopment of living tissue over time. OCM andMPM images from different days of the engineeredskin tissue under normal culturing conditions areshown in Figure 3. Comparison between the OCMand MPM images on different days further demon-strates the usefulness of the complementary informa-tion gained from the two modalities. It is seen fromthe time-lapse image sequence that more keratino-cytes have migrated from deeper layers to moresuperficial layers, and are losing nuclei and orga-nelles over time, which actually is one way to vali-date the process of keratinization. The process isfurther validated by the comparison of histology re-sults, which shows that the stratum corneum is thick-er after a few days. As shown by the images, theMPM images provide better visualization of the cel-lular structures during the early days, while theOCM images provide better observation of the kera-tinized cells during the later days of culture, wherecells are losing their normal morphology and orga-nelles.

Another important advantage of the multimodalmicroscope described here is the integration ofFLIM with the other two modalities, which extendsthe functional imaging capabilities of the system. Fig-ure 4 shows OCM and FLIM images of keratino-cytes at different epidermal layers. The FLIM imagesare composed of 256 � 256 pixels and were gener-ated in about 4 minutes. The lifetimes were analysedbased on a simulation model composed of two expo-nential components, which correspond to the freeand enzyme-bound NADH. The model is expressedby IðtÞ ¼ a1 e�t=t1 þ a2 e�t=t2, where a1, t2, and a2, t2

are the amplitudes and lifetimes of free and bound

NADH, respectively. The average lifetime is givenby tave ¼ ða1t1 þ a2t2Þ=ða1 þ a2Þ. It has been demon-strated that both the average lifetime and the ratioða1=a2Þ of amplitudes of the two components can beused to probe the metabolic activities of cells [39,41]. In this case, longer average lifetime or a smallerratio corresponds to a higher level of metabolic ac-tivity. The average lifetime or ratio shown in Fig-ure 4 was obtained by analysing the decay rate ofthe integrated fluorescence over the entire detectionspectrum of the FLIM system, which ranges from374 to 565 nm. Shown by the lifetime images and thecorresponding histograms in Figure 4, the averagelifetime is getting longer and the ratio of the twocomponents is getting lower, from the inner basallayer to the spinous layer. These results are in goodagreement with the difference in metabolic level ofkeratinocytes from different epidermal layers, i.e.,the lower metabolic states in the inner (basal) layersand higher metabolic states in the outer (spinous)layers.

It should be noted that the decay dynamic ofautofluorescence from a mixture of free and proteinbound NADH is a complex multi-exponential pro-cess, which may need a more complicated model anddata processing method to analyze [46]. Here it isapproximated using the frequently used approxima-

Figure 3 (online color at: www.biophotonics-journal.org)Time-lapse OCM and MPM images from an engineeredskin sample. (a) OCM (left column) and MPM (right col-umn) images from the stratum spinosum layers of the en-gineered skin after increasing days of culture. (b) Compar-ison of hematoxylin and eosin stained histology images ofthe engineered skin tissue on Day 1 (upper) and Day 7(lower) of culture. Scale bars represent 20 mm.


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tion of a bi-exponential decay model. In addition tothe level of the metabolism, the change in averagelifetime and/or ratio of the two exponential compo-nents is also an indicator of shifts between glycolyticand oxidative metabolism, which has been used forthe detection of a tumor margin [38]. This changecould also reflect the shorter lifetime (lower metabo-lism) of the basal layer within the engineered skin,compared to the in vivo skin, which could be attribut-ed to the lower oxygen concentration within the ba-sal layer of the engineered skin.

Incorporation of FLIM also adds new functionalcontrast capabilities to the imaging platform, whichis demonstrated in Figure 5, showing images of a mi-croporous fibroblast-seeded hydrogel scaffold thatwere obtained with different modalities. A centerwavelength of 760 nm was used for both modalities,and the incident laser power was 5 and 17 mW forOCM and FLIM, respectively. It is also shown inFigure 5 that the cell morphology and the micropor-ous structure of the hydrogel are clearly differen-tiated and well-observed with both OCM and MPM,in comparison to the phase contrast image. Becauseof the spectral overlap between the autofluorescencefrom the hydrogel and the autofluorescence fromcells, which is illustrated by spectroscopic analysis(Figure 5(b)), they are not well separated in theMPM image, which was generated based on thespectrally-integrated signals. In contrast, FLIMclearly distinguishes them based on the different life-times of the two fluorescence emissions, showing dif-ferent colors for the hydrogel and the cells. In thiscase, the lifetime was obtained by fitting the decaymeasurements with a single exponential model. Thespectra in Figure 5(b) were retrieved from the FLIMdata which was measured based on a 16-channelmonochromator.

OCM/FLIM images from different depths of invivo human skin are shown in Figure 6. The images

were obtained from a dorsal hand of a human volun-teer. The laser power used was 5 and 17 mW forOCM and FLIM, respectively, and the center wave-

Figure 4 (online color at:www.biophotonics-journal.org)OCM and FLIM images from anengineered skin tissue. (a) OCMimages from different depths (zeromicrons is at the surface) in theengineered skin tissue. (b) FLIMimages and corresponding histo-grams of average lifetime at differ-ent depths of the skin tissue. (c)FLIM images and correspondinghistograms of free-to-bound ampli-tude ratio at different depths ofthe skin tissue. (d) Pseudo-colorimages of OCM (top), FLIM(middle) and merged OCM andFLIM (bottom). Scale bars repre-sent 20 mm.

Figure 5 (online color at: www.biophotonics-journal.org)Multimodal images of a microporous 3D hydrogel scaffoldseeded with 3T3 fibroblasts. (a) Representative imagesbased on phase contrast, OCM, MPM, and FLIM. (b) Inte-grated spectra from two different areas in the FLIM imageshowing spectral overlap, but which can still be distinctly dif-ferentiated based on lifetimes. Scale bar represents 20 mm.




length was 760 nm for both the two modalities. EachFLIM image of in vivo skin was acquired in 30 s.FLIM images show the spatial distribution of aver-age lifetime which is based on the same proceduresfor generation of images in Figure 4(b). In vivo skinstructures at different depths, including the topogra-phy of the stratum corneum, the cellular morphology

of keratinocytes in the superficial epidermis, and thedermal papillae projections of the basal membranecan all be visualized in both OCM and FLIMimages, as shown in Figure 6. Lifetime changes (indi-cated by different colors) likely indicate alterationsin metabolic states and/or contributions from differ-ent chromophores. For example, the spectral analysisin Figure 6(b) shows strong SHG signal from theareas surrounding the papillae projections andbroadband autofluorescence from the cellular struc-tures. The fluorescence lifetime of in vivo skin issomewhat different from that of engineered skin.Specifically, the lifetime from the basal layer in invivo skin is longer than from the basal layer in theengineered skin, which could be due to the differ-ence in nutrient supply between in vivo natural skinand in vitro engineered skin. The measured lifetimefrom in vivo human skin in our study is in goodagreement with those in vivo FLIM results reportedby other researchers [30, 36]. Further correlations be-tween the measured data and the physiologicalchanges in the skin are needed, as well as more com-prehensive and systematic imaging investigations [36].

4. Discussion

As shown by the above results, the use of an inte-grated multimodal microscope offers the capabilityto visualize the complicated multilayer architectureof both in vitro engineered human skin and naturalin vivo human skin. The main structures of engi-neered skin, including fibroblasts and collagen in thedermis, and different morphologies and physiologicalstates of keratinocytes in the epidermis, are visua-lized by depth-resolved images. Surface topography,cellular morphology, and 3D projections of dermalpapillae are observed in in vivo human skin imaging.Complementary information obtained with this inte-grated microscope consists of the structural architec-ture, cellular morphology, and metabolic states,based on the corresponding OCM, MPM, and FLIMmodalities. The spatially co-registered images fromdifferent modalities actually provide a matrix of vec-tors for each pixel, rather than the one dimensionalscalar information provided by any single modality.We believe this matrix information will enable oneto ultimately obtain a better understanding of thephysiology of the localized tissue or cells.

Time-lapse images demonstrate the capability ofthis microscope and methodology for monitoring thedevelopmental processes of skin in real time. Thesefeatures enable one to investigate the developmentand cellular dynamics in engineered skin tissues, aswell as facilitate our understanding of the functionand physiology of natural in vivo human skin. Forexample, the first and most important protective bar-

Figure 6 (online color at: www.biophotonics-journal.org)OCM and FLIM images of in vivo human skin. (a) OCM(left column) and FLIM (right colum) images. (b) Inte-grated FLIM spectra of two different areas noted in theFLIM image from the basal membrane layer. Abbrevia-tions: SC, stratum corneum; ED, epidermis; SB, stratumbasale; BM, basal membrane; D, dermis. Scale bar repre-sents 20 mm.


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rier of skin is guaranteed by the characteristic archi-tecture of the stratum corneum. This uppermostlayer of the epidermis consists of dead cells (corneo-cytes) that lack nuclei and organelles. Generationand renewal of the stratum corneum is ensured bythe transformation of keratinocytes into corneocytes[47]. This process, which is the main feature of epi-dermal differentiation, starts with the proliferationof keratinocytes in the stratum basale and ends withthe cornification of cells (technically dead). Duringthis process, the keratinocytes migrate toward thesurface through the epidermal layers, including thestratum basale, stratum spinosum, stratum granulo-sum, and the stratum corneum. The cells appear withdifferent morphological and physiological states witheach of these layers. As shown by the images in Fig-ures 2 and 4, the morphological differences in theseepidermal layers are readily revealed by OCM andMPM images, and the different metabolic states as-sociated with these different physiological states aremanifested in the FLIM images. Time-lapse imagesin Figure 3 also demonstrate the keratinization pro-cess over time.

While each modality has unique features for im-aging the tissue, the complementary information pro-vided by the spatially co-registered images providesadditional insight into the tissue structure and func-tion, and mutually validates the results obtained withthe individual modalities. As shown by the OCMand MPM images in Figure 3, while MPM providesbetter visualization of the cellular structure, OCMshows advantages when cells are close to cornifica-tion and lose their cellular shape and organelles(and thus the spatially confined fluorophores onwhich MPM images are based). In practical imagingprocedures, the different modalities also facilitateeach other to obtain the desired imaging results. Forexample, because OCM is advantageous for imagingthe overall structures present within the samplebased on the scattering contrast, and it uses rela-tively lower incident laser power, OCM is alwaysused as the guiding mode to locate the desired im-aging area of interest. This is particularly important inthe case of using the OCM and FLIM modes, giventhe relatively slow imaging speed of FLIM (usuallyseveral minutes for one image). The slow imagingspeed of FLIM also necessitates the use of the fastMPM mode for locating desired regions of interest,although MPM images can also be generated basedon the FLIM data. The complicated processes forsignal detection in FLIM, including fiber collectionof fluorescence signal, spectral separation by themonochromator, and time-resolved photon countingapproaches, result in the low signal detection effi-ciency and thus the slow imaging speed for FLIM.Therefore, the MPM mode with much higher signaldetection efficiency and imaging speed is more oftenthe preferred way for generating MPM images.

OCM offers other potential capabilities for theinvestigation of engineered and natural skin. Com-pared to reflection mode confocal microscopy, whichis another widely used microscopy technique that isable to generate 3D images based on the scatteringproperties of samples, OCM is advantageous in thatits axial resolution is determined by the bandwidthof light source rather than the numerical aperture ofthe objective. This feature enables the OCM modal-ity to be flexibly modified for a larger field-of-viewand depth-of-field. In our system, this can be donebe changing the beam diameter of the OCM beamat the back aperture of the objective to alter the ef-fective NA of OCM while keeping the MPM param-eters unaffected. Because OCM utilizes coherencegating in addition to the confocal gating utilized inconfocal microscopy, OCM is able to image consider-ably deeper into high-scattering tissues, such as skin,because the added coherence gating rejects multiply-scattered photons that only contribute to back-ground noise in confocal microscopy [48]. There arealso many variants of OCT/OCM that allow comple-mentary features of the sample to be observed.These variants include polarization-sensitive OCTfor visualizing tissue birefringence due to the pre-sence and orientation of collagen fibers [49, 50], op-tical coherence elastography that is capable of meas-uring biomechanical parameters of the sample [51–53], and magnetomotive OCT used to enhance im-aging contrast [54]. These additive imaging modalitieswill be explored in the future to further characterizethe structural morphology and the functional andmolecular physiology of engineered and naturalskin.

5. Conclusions

In conclusion, we have demonstrated a unique inte-grated OCM/MPM/FLIM microscope that is capableof multimodal characterization and evaluation of en-gineered and natural skin tissues. With the capabilityfor in vivo imaging of human skin, this multimodalmicroscope also represents a potentially powerfultool for clinical dermatological studies, provided thatremaining technical issues are addressed, such as abetter design of a dark box or a miniaturized straylight-limiting enclosure surrounding the imaging site,and increased flexibility of the imaging probe withbetter optical arrangement. Combined with micro-scope-compatible bioreactors, which are under de-sign development in our lab, this imaging platformwill enable one to monitor the developmental pro-cesses of engineered skin tissue in culture, and thecellular and functional dynamics of skin tissues dur-ing processes such as repair and regeneration inwound healing. These results lay the foundation for




future work using this multimodal microscopy ap-proach longitudinally over time to investigate andoptimize the culturing of engineered skin tissues, totrack cell and tissue changes under different bio-chemical and biomechanical culturing conditions,and to evaluate the in vivo host skin response duringthe engraftment of engineered skin used in the treat-ment for skin injury or loss.

Acknowledgements We thank Dr. Marina Marjanovicfor her helpful discussions and assistance with this re-search, and Mr. Darold Spillman for his technical, logisti-cal, and information-technology support to this research.This research was supported in part by grants from theNational Science Foundation (CBET 08-52658 ARRA,CBET 10-33906, S.A.B.). Benedikt Graf was supported bythe Pre-doctoral National Institutes of Health Environ-mental Health Sciences Training Program in Endocrine,Developmental and Reproductive Toxicology at the Uni-versity of Illinois at Urbana-Champaign. Additional infor-mation can be found at http://biophotonics.illinois.edu.

Youbo Zhao received his B.S. degree in applied physicsfrom Beijing Institute of Technology, Beijing, China, in1997, and his Ph.D. degree in optical engineering fromNankai University, Tianjin, China, in 2007. He is cur-rently a postdoctoral research associate in the Biopho-tonics Imaging Laboratory at the University of Illinoisat Urbana-Champaign, Urbana, IL. His research inter-ests include developing and using novel optical imagingtechnologies for biomedical studies.

Benedikt W. Graf is a Ph.D. candidate in electrical andcomputer engineering at the University of Illinois atUrbana-Champaign. He is a research assistant in theBiophotonics Imaging Laboratory at the Beckman In-stitute for Advanced Science and Technology. His re-search interests include the development of multimodaloptical imaging techniques for in vivo imaging of skinfor both clinical and research applications. Researchapplications include tracking of stem cells in live mouseskin during wound healing.

Eric J. Chaney received his B.S. degree in biology fromthe University of Evansville, Evansville, Indiana, in1992. From 1993 to 1997, he worked as a research assist-ant at the Indiana University School of Medicine, Indi-ana State University, Terre Haute. From 1997 to 2000,he worked as a transmission electron microscope tech-nician at the University of Illinois at Urbana-Cham-paign. Since 2000, he has been a research scientist inmolecular biology at the Biophotonics Imaging Labora-tory, Beckman Institute, University of Illinois, Urbana-Champaign.

Ziad Mahmassani received his B.S. degree in bioengi-neering from the University of Maryland, College Park,Maryland. He is now a graduate student in the Depart-ment of Kinesiology and Community Health at theUniversity of Illinois at Urbana-Champaign. His re-search focuses on regulation of muscle stem cells bymechanical strain.

Eleni V. Antoniadou is a Ph.D. candidate in bioengi-neering at the University of Illinois at Urbana-Cham-paign, USA. She holds a M.S. degree in nanotechnol-ogy and regenerative medicine from the MedicalSchool of the University College London. She is cur-rently a research assistant in the Tissue Engineering,Biomaterials and Stem Cell Niche Lab of the Chemicaland Biomolecular Engineering Department. Her re-search interests include regenerative medicine applica-tions, the development of artificial organs, and tissueengineering.

Ross DeVolder received his B.S. degree in ChemicalEngineering from the University of Iowa in 2008, andhis M.S. in Chemical Engineering from the Universityof Illinois in 2010. He is currently a graduate student inthe chemical engineering department at the Universityof Illinois, Urbana-Champaign, working in Dr. Hyun-joon Kong’s research group. His research interests in-clude designing polymeric based materials with control-lable properties for a wide array of applications.

Hyunjoon Kong is an assistant professor in the Depart-ment of Chemical and Biomolecular Engineering. Healso holds affiliations with the Department of Bioengi-neering, Center for Biophysics, and ComputationalBiology and Neuroscience Program. He received hisengineering education from the University of Michigan(Ph.D. 2001), and performed post-doctoral research atthe University of Michigan and Harvard University. Hejoined the University of Illinois in 2007. His current re-search is focused on the design of biomaterials used fortissue regeneration, drug delivery, and imaging.

Marni D. Boppart is an assistant professor in the Col-lege of Applied Health Sciences, University of Illinoisat Urbana-Champaign (UIUC), and holds a full-timefaculty appointment at the Beckman Institute. She re-ceived her B.S. in Cell Biology from the University ofNew Hampshire, her M.S. in Cell Biology fromCreighton University, and her Sc.D. in Applied Anato-my and Physiology from Boston University. She com-pleted postdoctoral training at Harvard University andUIUC in Cell and Developmental Biology. Her currentresearch interests include cellular biomechanics and therole for mesenchymal stem cells in muscle.


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BIOPHOTONICS


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Stephen A. Boppart is currently a Bliss Professor ofEngineering in the Departments of Electrical and Com-puter Engineering, Bioengineering, and Medicine, Uni-versity of Illinois at Urbana-Champaign. He receivedthe B.S. degree in electrical and bioengineering and theM.S. degree in electrical engineering from the Univer-sity of Illinois at Urbana-Champaign, Urbana, IL, in1990 and 1991, respectively, the Ph.D. degree in electri-cal and medical engineering from the Massachusetts In-stitute of Technology, Cambridge, in 1998, and theM.D. degree from Harvard Medical School, Boston,MA, in 2000. His research interests include the devel-opment of novel optical imaging technologies for biolo-gical and medical applications, with particular emphasison translating these to clinical applications.




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