Single laser source for multimodal coherentanti-Stokes Raman scattering
microscopy
Adrian F. Pegoraro,1,2,* Aaron D. Slepkov,2 Andrew Ridsdale,2
John Paul Pezacki,2 and Albert Stolow1,2
1Department of Physics, Queen’s University, Kingston, Ontario K7L 3N6, Canada2Steacie Institute for Molecular Sciences, National Research Council of Canada,
Laser use continues to make major inroads into bio-medical research, with confocal laser scanningmicro-scopy (CLSM) providing a prime example [1]. Themost common CLSM systems use fluorescence forimaging, wherein the excitation wavelength is chan-ged to measure different fluorescent dyes and mole-cules. The introduction of nonlinear optical (NLO)techniques has opened new doors for CLSM. For ex-ample, it has led both to new light sources for fluor-escence microscopy [2] and to new techniques, suchas stimulated emission depletion microscopy, to im-prove spatial resolution [3]. Pulsed lasers also enablenew imaging modalities for CLSM. The most widelyutilized NLO modality is two-photon fluorescence(TPF) imaging [4]. Other NLO techniques have alsobeen used for imaging, such as second harmonic gen-eration (SHG) [5] and third harmonic generation(THG) imaging [6]. NLO microscopy can allow con-
trast without requiring the addition of labels or dyes.Tissue samples, in particular, often demonstratestrong intrinsic nonlinear signals. In multimodalimaging, TPF, SHG, and THG make use of a singlepump wavelength, using color filters to separate thesignals; however, these do not demonstrate chemicalspecificity. For chemical specific imaging, confocalRaman is a well known technique, but it requireslong acquisition times that are generally incompati-ble with live-cell microscopy. Coherent anti-StokesRaman scattering (CARS) microscopy is a NLOimaging modality that overcomes this deficiency byusing two lasers with their frequency differencetuned to a Raman vibrational resonance [7,8]. It isthis driven (stimulated) aspect of CARS that leadsto its greatly enhanced signal levels in comparisonwith spontaneous Raman scattering and enablesfaster imaging. By changing the wavelength of oneor both lasers, different Raman resonances can beprobed to better characterize the sample. CARS thusallows label-free, chemical-specific imaging and isbecoming an important tool for the study of livesamples. Here we discuss a simplified approach to
multimodal CARS microscopy that we hope willencourage its expanded use in biomedical research.
Although CARS does offer significant advantagesover other NLO modalities, there are also challengesthat must be overcome for effective imaging, themost significant being the nonresonant background.The energy level diagrams for resonant and non-resonant CARS are illustrated in Fig. 1. We can seethat both the resonant and the nonresonant compo-nents are indistinguishable by signal wavelength.Although this is not an issue for bulk samples com-posed only of resonant material, inmicroscopy, wherethe concentration of resonant molecules is small, thispresents significant difficulties. There are manytechniques that aim to overcome the nonresonantbackground and several review articles are availablethat highlight these [9–12]. The most common meth-od of overcoming this problem is to use picosecond(ps) instead of femtosecond (fs) pulses to improvethe contrast of resonant over nonresonant signals[13]. This has traditionally limited the choice of lasersource commonly used for CARSmicroscopy to eithertwo synchronized ps oscillators [14] or to a ps opticalparametric oscillator (OPO) [15]. One drawback ofthis approach is that the use of ps pulses for CARSmicroscopy makes it more difficult to combine CARSwith other NLO techniques in a multimodal micro-scope, since ultrashort (fs) pulsed lasers are the cur-rent tool of choice for NLO imaging. Although it ispossible to acquire CARS and other NLO signals se-quentially, ideally all signals should be acquired si-multaneously. The question of which laser source touse, of course, depends on the application. The bestsolution for working in the fingerprint region(Raman shifts<1700 cm−1) is not necessarily the bestchoice for lipid imaging (Raman shifts from 2700 to3100 cm−1). To clarify this discussion, in Section 2 weshow how the pulse spectral width is related to CARSmicroscopy performance, both in imaging and micro-spectroscopy applications. This highlights the needfor flexible laser sources in multimodal CARS micro-scopy. We present a discussion of our simplified, fs-based approach to CARS microscopy that recently
permitted commercialization of the first CARS mi-croscope specifically designed for biomedical imagingof lipids [16].
2. Theory
Both ps and fs pulses can be used for CARSmultimo-dal microscopy. The ideal laser for CARS microscopydepends on many factors, the most important onebeing the Raman vibration(s) probed. This is best il-lustrated with a simple model simulation, followingthe method of Cheng et al.[13]. Here we compare theperformance of different pulse spectral widths for agiven sample. The CARS polarization PCARSðωasÞ isgiven by
PCARSðωasÞ ¼Z þ∞
−∞
χð3ÞEpðωpÞE�sðωsÞEprðωprÞdω; ð1Þ
where EpðωpÞ, EsðωsÞ, and EprðωprÞ are the electricfields of the pump, Stokes, and probe pulses, respec-tively. For degenerate CARS microscopy, the pumpand probe are identical. The third-order susceptibil-ity is given by χð3Þ ¼ χð3Þr þ χð3Þnr , where χð3Þr and χð3Þnr de-note the resonant and nonresonant susceptibilities,respectively. To model a single Raman resonancewe use
χð3Þr ¼ AΩ − ðωp − ωsÞ − iΓ ; ð2Þ
where Ω is the center frequency of the Raman mode,Γ is the linewidth, and A is a scale factor. The laserpulses are normalized to be of constant energy, re-gardless of spectral width. This choice of normaliza-tion has further implications when photodamage istaken into consideration, as discussed below:
EðtÞ ¼ E
ffiffiffiffiffiffiffiffiffiffiffiffiπδ
2 ln 2
rexp
�−π2δ2t22 ln 2
�expð−iωtÞ; ð3Þ
whereE is amplitude, δ is the spectral width, and ω isthe center frequency of the pulse.
To enable comparisons for practical imaging appli-cations,weuseparameterssuitable for imaging lipids:a pumpcenter frequency of12500 cm−1 (800nm)andaStokes center frequency variable between 9575 and9700 cm−1 (1030–1044nm). The energy difference be-tween the two pulses covers the 2800–2925 cm−1
range.Weuse twooverlapped resonantRamanmodestosimulateacomplexRamanstructure.Onemodehasa Raman shift Ω1 of 2850 cm−1 and a linewidth of2Γ1 ¼ 15 cm−1; the other has a resonant frequency ofΩ2 of 2875 cm−1 and a Raman linewidth of 2Γ2 ¼20 cm−1. The mode centered at 2875 cm−1 is twice asstrong as the one centered at 2850 cm−1 (A2875 ¼2A2850). The relative strength of the resonant to non-resonant signals modeled is jA2850=χð3Þnr j ¼ 0:01 cm−1
(this approximates the contrast achieved in practicalsystems [17]). For ease of comparison, the intensity
Fig. 1. (a) Energy level diagram for resonant CARS. The energydifference between pump ωp and Stokes ωs matches the energy ofthe Raman level to be probed ΩR. The anti-Stokes ωas light is thedesired signal. (b) Energy level diagrams for nonresonant CARS.The anti-Stokes light generated has the same frequency as isgenerated in the resonant case, yet does not rely on ground-statevibrational coherence.
from the resonant and nonresonant contributions toχð3Þ is calculated separately.
In Fig. 2(a), the calculated resonant signals (inten-sity Ir) are shown as a function of Stokes centerwavelength for different pulse durations. As ex-pected, total signals increase and spectral resolutiondecreases as the spectral width of the pulses is in-creased (i.e., toward fs pulses). Figure 2(b) showsthat the contrast (resonant divided by nonresonantintensity Ir=Inr) improves with a smaller spectralwidth (i.e., toward ps pulses). This strongly suggestsan advantage to the use of spectrally narrow pulsesand is one of the primary reasons for choosing pspulses; however, signal levels are also sacrificed withthis choice. To characterize the compromise betweencontrast and desired resonant signal, we measurethe performance Pf , which is defined as the contrastmultiplied by the resonant signal (Pf ¼ I2r=Inr). Theperformance as a function of Stokes center wave-length for different pulse spectral widths is shownin Fig. 2(c).
Figure 2 illustrates some of the considerations fordetermining which pulse spectral width to use for agiven sample. In this example, performance is max-imized for pulses with a spectral width of ∼35 cm−1
and peaks at a frequency difference of ∼2870 cm−1. Ifthe combined Raman vibrations are approximated asa single Raman mode, this corresponds to the effec-tive center frequency and Raman bandwidth. Thisdemonstrates that for CARS imaging applicationswhere it may not be critical to distinguish betweenadjacent Raman modes (e.g., lipid imaging in tis-sues), maximum performance is achieved by use of apulse spectral width that matches the effective spec-tral width of the overlapped Raman modes, not thewidth of any particular band. For samples in whichlipid imaging is used, this corresponds to pulses witha spectral width >100 cm−1 (<150 fs). On the otherhand, if it is necessary to perform imaging while iso-lating a single Raman band, pulses that are spec-trally narrow relative to the Raman resonanceunder observation are required, since they enable
contrast between adjacent bands. This is particularlyrelevant in the fingerprint region in which there aremany overlapped Raman bands with linewidths<10 cm−1 (>1:5ps), which can produce a large nonre-sonant background when trying to image a specificband. This implies that a laser source with variablepulse spectral width is preferred for multimodalCARS imaging. The desire for flexibility in choosingthe pulse spectral width is further reinforced whenconsidering photodamage in NLO microscopy. If themechanism of photodamage is a higher-order non-linear process than the process that generates thesignal, peak power could be used for normalizationrather than constant energy as was employed above.In this case, to maintain a constant output signal, itwould be preferable to use longer pulses with a high-er average power. Eventually photodamage that isdue to absorption in the sample would limit increas-ing the average power and thus set the optimal pulseduration. For TPF, the primary damage mechanismwas found to scale linearly [18], quadratically [19], orfaster than quadratically with intensity dependingon the sample and laser used [20]. For CARS micro-scopy, photodamage has been shown to scale betweenlinearly and quadratically with intensity dependingon the sample and scanning conditions [21]. Conse-quently, peak power can be the sample-dependentlimiting factor that favors longer pulses whereas theopposite is true when limited by average power. Thisagain highlights the desire for flexibility in choosingthe laser source used. Varying the pulse spectralwidth to either maximize performance or minimizephotodamage can be accomplished by using multiplelaser sources [22], using long doubling crystals tonarrow the spectral bandwidth [23] or by temporallystretching (chirping) spectrally broad pulses toachieve an effectively narrower spectral bandwidth[17,24–27]. This last option is known as spectral fo-cusing and allows the effective pulse spectral widthat second order (the Raman resonance) to be continu-ously adjusted to the desired effective pulse band-width. This offers the further advantage that
Fig. 2. The CARS (a) resonant signal, (b) contrast, and (c) performance are shown as a function of pump-Stokes frequency difference forvarying pulse spectral widths for pulses of constant energy. The broadest (i.e., temporally shortest) pulse achieves the maximum resonantsignal whereas maximum contrast is achieved for the narrowest (i.e., temporally longest) pulse. The individual peaks of the Raman spec-trum are resolved only by use of the narrowest pulse. Performance is maximized for a pulse spectral width that matches the effectivelinewidth of the combined Raman mode; see text for definitions of contrast and performance.
optimization is performed after generation (i.e., itdoes not require modifying the laser) and is easily ad-justed to suit a particular sample. Finally, this allowsreal-time feedback optimization of the CARS spectralresolution, image contrast, as well as signal strengthfrom other NLO signals, as the effective pulse widthis varied.
In microspectroscopy applications, it is the Ramanspectrum of the sample that is measured. A straight-forward approach to extracting the Raman spectrumusing a CARS microscope is to use CARS imaging tolocate points of interest in a sample and then toswitch to confocal Raman microscopy to extract thespectral information [28]. This solution is particu-larly attractive for use in the fingerprint region, sinceit has been shown that confocal Raman offers com-parable sensitivity to CARS in this region for live-cellmicroscopy applications [29]. If the CARS signalitself is used to extract spectral information, theprocedure is more complex, but techniques are avail-able. Multiplex CARS uses a ps pump and fs Stokesto probe many Ramanmodes simultaneously and theCARS spectrum is collected using a spectrometer[30,31]. Because of interference between the nonreso-nant background and the resonant signals, postpro-cessing of the CARS spectrum is required to extractthe Raman spectrum [32]. More sophisticated techni-ques utilizing interference within fs pulses are alsoavailable [33,34]. A more direct approach is to recordthe CARS spectrum by varying the wavelength of onelaser and reimaging the sample; however, interfer-ence with the nonresonant background causes theCARS and Raman spectra not to be identical. Eitherpostprocessing or using interferometric techniquesto retrieve the Raman spectrum can overcome thisdifficulty [35]. Regardless of the method used to ex-tract the Raman spectrum, the effective pulse spec-tral width limits the achievable spectral resolution.It is worth noting that, when using chirped fs pulses,rather than changing the laser center wavelength itis possible to scan the frequency difference (i.e., theRaman resonance being probed) simply by changingthe time delay between the pulses [17,26]. This is asimple way to perform spectral scans without adjust-ing the laser source and enables much faster spectralscanning. Again, the spectral resolution is optimizedby changing the pulse chirp rate via dispersionengineering.
3. Toward All-Fiber Lasers
Regardless of the chosenpulsewidth, systemstabilityand ease of use are primary concerns of microscopyendusers. InCARSmicroscopes that use two synchro-nized lasers, timing jitter and alignment duringspectral tuning are problematic and put additionalconstraints on the operating environment (tempera-ture and vibrational stability) [14]. Use of an OPOto generate the pump and Stokes removes the pro-blems with timing jitter and alignment during spec-tral tuning [15,27].Use of a single laser andaphotoniccrystal fiber (PCF) to generate the Stokes, however,
automatically synchronizes the pulses and repre-sents a simple and cost-efficient method for perform-ing CARS microscopy [17,36,37]. Laser alignmentinto the fiber is prone to vibrations and environmen-tal changes, but automated fiber alignment feedbacksystems exist that can compensate for these drifts.These solutions rely on free-space lasers, whereasan all-fiber solution should offer improved stabilityand alleviate concerns about system alignment. Italso offers a natural path to future integration withan endoscope for in vivo imaging. This was recognizedearly on inCARSmicroscopy and effortsweremade toinclude fiber components for delivery of the excitationbeams [38]. The main challenge using all-fiber gen-eration for CARS microscopy is to generate a second,synchronized laser pulse at a different center wave-length. Replacing the pump laser with a fiber laserin anOPOapparatus is one approach to incorporatingfiber lasers into CARS microscopy [39] but stillrequires generating the Stokes using free-spacecomponents.
Soliton shifting in highly nonlinear fibers plays akey role in generating the pulse pair in generation ap-proaches [23,40,41]. Regardless of the nonlinear fiberused, the output pulses tend to be spectrally broad(>200 cm−1) and must be accommodated if high spec-tral resolution is desired. Another concern is that thesoliton generation process in fibers can be unstableunless the input pulse to the fiber is on the short fstime scale and transform limited [42,43]. This pre-sents challenges for CARS microscopy where, in gen-eral, ps pulses have beenpreferred.One approach hasbeen to design a system formultiplex CARS but to sa-crifice rapid imaging, which is done with a ps fiberpump laser and then the use of two cascaded non-linear fibers to first broaden then soliton shift the out-put to generate the Stokes pulse [40]. Use of a fs fiberlaser alleviates the need to broaden the pulse beforeStokes generation, but the resulting broad bandwidthpulse must be compensated by use of another techni-que if better spectral resolution is required for ima-ging. One method is to spectrally narrow both thepump and the Stokes by using long periodically poleddoubling crystals [23]. On the other hand, use of spec-tral focusing to achieve the desired bandwidth offersthe advantage of being able to easily tune the effectivepulse spectralwidthwhile enabling otherNLOmicro-scopies by the short pulse [41]. Currently, all these ap-proaches still use some free-space components andhave relatively lowpower in comparisonwith existingfree-space lasers; however, they present promisingideas for further development of stable, high averagepower, fiber laser sources for multimodal CARSmicroscopy.
4. Microscopy Setup
A depiction of a general CARS microscope opticalarrangement is given in Fig. 3. A pump laser is com-bined with a method to generate the Stokes pulse.The Stokes can be generated by a second synchro-nized laser [14], an OPO [15,27], or a PCF [17,36,37]
as discussed above. The power utilized is limited bysample damage, with typical power levels of the or-der of 10–50mW for live-cell imaging (assuming∼80MHz pulse repetition rate and a pixel dwell timeof ∼4–8 μs). Integrated laser exposure time plays asignificant role in determining cell damage, so in-creasing the microscope scan speed (i.e., reducing thepixel dwell time) allows for higher laser power at thesample [44]. Using wavelengths closer to the infraredalso reduces sample damage [19]; however, waterabsorption sets an upper limit on the Stokes wave-length (at ∼1:5 μm). Longer wavelengths also offerthe added advantage of reducing scattering lossesthat enable greater penetration depth into samples.This has been most successfully demonstrated withps OPOs with a degeneracy point at 1:05 μm, whichprovides pump and Stokes wavelengths at ∼900 and1260nm, respectively; however, if comparable fssources are developed, they should provide the samebenefits. The microscope used for CARS imaging canvary considerably depending on the application.Most systems make use of objectives with a highnumerical aperture (NA) that allow the pump andStokes to be sent into the microscope collinearly.Tight focusing relaxes the phase-matching conditiontypically requisite for a four-wave mixing processsuch as CARS. The CARS signal is much strongerin the forward direction, and this is often used ifthe sample permits it [13] . For thick tissue samplesand live animal studies, the signal must of course becollected in the epidirection. For multimodal ima-ging, the different signals are separated by use ofdichroic beam splitters and filters.
5. Applications
Here we present applications of multimodal CARSmicroscopy, including those based on the simplifiedapproach highlighted in this article. Much of the in-terest in optimizing the source for multimodal CARSmicroscopy derives from the advantages it offers forthe study of live cells and tissue. Although CARScan be used to image different Raman modes (e.g.,D2O to study water dynamics [45]), most applicationsfocus on lipid imaging using CARS, due to its strong,
relatively isolated Raman band and high concentra-tion of oscillators. When complemented with othermicroscopies, CARS imaging provides new informa-tion about the sample. One technique that can be ad-vantageously combined with CARS is differentialinterference contrast (DIC), which offers label-freeimages primarily based on density changes in thesample. DIC can be enhanced by using CARS to high-light objects of interest based on their Raman signa-ture [46]. Because DIC is a low power technique, thisallows rapid imaging over long time periods withoutfear of cellular damage. We, for example, used thiscombination to study the effects of the Hepatitis Cvirus (HCV) on liver cells.HCV is known to cause lipiddrop accumulation [47,48]; to study these changesCARS is used to locate lipid drops whereas DIC isused to track them. An example showing a CARS im-age taken using chirped fs pulses overlaid on DIC isshown in Fig. 4(a). CARS can also be combined withtraditional fluorescent labeling techniques to offer en-hanced information. If CARS and TPF are measuredusing the same laser, it ensures that the signals arecoregistered and allows for collocalization studies.This has been used to study several different pro-blems in cellular dynamics. Figure 4(b) again high-lights the use of CARS to study HCV. Here, the HCVconstruct has been modified to induce the productionof green fluorescent protein (GFP). TPF is used toidentify infected cells, and CARS is used to track lipidmovement and accumulation. Chirped fs pulses areused to collect both signals simultaneously.
Tissue samples offer another area in which multi-modal NLO microscopy provides significant advan-tages.Although fluorescent staining canbe employed,many samples demonstrate endogenous fluorescencethat obviates the need for labeling. Furthermore, the
Fig. 3. Generalized CARS microscopy optical setup. The pumplaser is split using a beam splitter (BS). Part of the beam is usedto generate a Stokes pulse; the other part is delayed then recom-bined and overlapped with the Stokes on a dichroic mirror (DM).Inside the microscope galvo-scanning mirrors (GS) direct thebeam. Dichroic mirrors and bandpass (BP) filters separate the sig-nals from the excitation light. Epidetected signals can be collectedbefore or after the galvo mirrors (descanned or nondescanneddetection).
Fig. 4. (Color online) (a) CARS and DIC multimodal imaging ofHuH-7 liver cells. The CARS image was taken first using chirped fspulses and then overlapped with a subsequent DIC image. Inprint, some of the lipid drops are indicated; online, red representsCARS due to lipids, gray represents the DIC image. Scale bar is5 μm. (b) CARS and TPFmultimodal image of HuH-7 cells infectedwith a modified HCV strain (infected cells express GFP). Theimages were acquired simultaneously using chirped fs pulses aspart of a time course studying live cells (taken at the 48 h timepoint). In print, the areas indicated are TPF due to GFP; the re-mainder of the image is CARS due to lipids. Online, greenrepre-sents TPF due to GFP; red represents CARS due to lipids.Scale bar is 10 μm.
long-range structure present in tissues allows SHG tobeusedasathirdcontrastmechanism.Sumfrequencygeneration and THG can also be used, offering addi-tional tools for the study of samples. Of particular in-terest is combining CARS, SHG and TPF since allthree signals can be collected using the same lasersand are often spectrally distinct. This has been de-monstrated for the study of atherosclerotic plaquein which all three signals have endogenous sources[22]. An example image taken using chirped pulsesis showninFig.5(a),where theblue isdue toSHGfromcollagen that composes the cap of the arterial plaque,green is TPF fromelastin in the arterywall, and red isCARS from lipids located inside the plaque. This par-ticular sample is from a Watanabe heritable hyperli-pidemic myocardial infarction (WHHLMI) rabbitstrain [49] where the disease progression is wellknown. By characterizing plaques using traditionaltechniques as well as NLO microscopy, it is hoped todesignmethods of diagnosing plaque severity in vivo.Thin tissue samples offer the additional advantage ofallowing collection in both the forward and the back-ward direction, which enables studying the ratio of aforward to backward collected signal to better charac-terize tissue composition. This is important in under-standing disease as well as basic tissue morphology.For example, by using SHG, CARS, and THG, it hasbeendemonstratedthatordering intissuecanbechar-acterized as the techniques are sensitive to differentheterogeneities [50]. For example, SHG [Fig. 5(b)],CARS [Fig. 5(c)], and THG [Fig. 5(d)] all reveal differ-ent information about fascia samples (TPF and THGtaken using fs pulses, CARS using ps pulses).Although SHG measures the heterogeneity of χð2Þ,THG is used to measure interfaces between collagensheets and CARS is used to measure bulk structure.The results suggest that SHG in collagen is due to im-perfect cancellation rather than from individual col-lagen fibrils, since THG and CARS do not indicatethe same level of heterogeneity.
Multimodal NLO microscopy also applies to livetissue imaging in which endogenous signals can beutilized without damaging the sample. The two mainalterations from working with cells and thin tissuesamples are that all the signals must be collectedin the epidirection and that long working distance
microscope objectives are preferred. Both CARS andSHG are emitted in the epidirection if the particles ofinterest are subwavelength; however, the forwardpropagating signal also rescatters in the backwarddirection and represents the majority of the signalcollected [44,51]. Given the large lipid content inmyelin, CARS has been particularly successful instudying nerve damage and recovery [52–54]. Whenimaging in live animals, longitudinal studies arepossible to better understand damage to the spine,coping mechanisms, and long-term recovery. Anotherexample is to use CARS to study white matter in the
Fig. 5. (Color online) (a) WHHMLI arterial sample. Online, redrepresentsCARS imagingof lipids, green representsTPF fromelas-tin, and blue represents SHG from collagen. In print, the differentparts of the sample are labeled.All the signalswere collected simul-taneously using chirped fs pulses. Scale bar is 50 μm. Multimodalnonlinear optical imaging of fascia. (b) Forward SHG image usingfs pulses, (c) forward CARS image at Raman shift frequency Δω ¼2845 cm−1 using ps pulses, and (d) forward THG image using fspulses. Scanning area was 146:2 μm2 (512 × 512 pixels). (b), (c)and (d) reprinted with permission from [50].
Fig. 6. In vivo CARS imaging using ps pulses of mouse brain using an upright microscope with a dipping mode water objective. (a) Aschematic of the in vivo CARS microscope; D, dichroic mirror. (b) CARS image of the parietal cortex. (c) Epi-CARS image of bundles ofmyelinated fibers in the subcortex white matter. Reprinted with permission from [52].
brain, as shown in Fig. 6 (ps pulses were used) [52].Currently, most live tissue work makes use of micro-surgery to expose the area of interest, highlightingthe need for endoscopic options for multimodal mi-croscopy to increase its utility and versatility.
While most applications of CARS microscopy havefocused on imaging a single Raman resonance, ex-tracting spectral information is used in some appli-cations. For example, multiplex CARS has been usedto study the effects of temperature change on lipiddomains [32]. CARS combined with confocal Ramanhas been used to study the differences between lipiddroplets in live animal versus cell culturemodels, im-portant for understanding and validating cell models[28]. This could be used, for example, to study the ef-fect of different obesity treatments from drug use tothe effects of dietary changes.
6. Conclusion and Outlook
Laser technology now plays a key role in biomedicalresearch, and we believe that multimodal NLO opti-cal microscopy, in particular, will continue to play anexpanding role in the life sciences. Label-free tech-niques are particularly important for studying livesamples in both cells and animal models. CARS mi-croscopy offers a new label-free modality that playsan increasingly important role in multimodal micro-scopy. Development in laser sources will be a key en-abling technology for the spread of this technique.Fiber laserswill likely play an increasingly importantrole in NLO microscopy, enabling use in a wider vari-ety of settings. Using fiber lasers will also make cou-pling to endoscopes substantially easier, whichwill benecessary if NLO microscopy is to be used for patientdiagnostics (beyond the skin and eye) without requir-ing surgery for imaging.Existing sources, however, al-ready allow effective multimodal CARS microscopyfor biomedical research. As we have shown here, sim-plified sources that offer the user control over the ef-fective pulse spectral bandwidth are versatile formultimodal NLOmicroscopy. Recently, this approachwas successfully implemented by a major microscopemanufacturer (Olympus America) to produce theworld’s first commercially available CARS micro-scope [16]. These developments along with the workof many other research groups will continue to makemultimodal CARS microscopy available to study awide variety of problems in biomedical research.
The assistance of Rodney Lyn for the HuH-7 cellsand Alex Ko at the Institute for Biodiagnostics,National Research Council of Canada, Winnipeg,Manitoba, Canada, for rabbit artery samples aregratefully acknowledged. A. Pegoraro thanks theNational Sciences and Engineering Research Coun-cil of Canada (NSERC) and the National ResearchCouncil of Canada (NRC) for funding support.
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