532 OPTICS LETTERS / Vol. 35, No. 4 / February 15, 2010

Multiphoton coherence domain molecular imagingwith pump–probe optical coherence microscopy

Qiujie Wan and Brian E. Applegate*Department of Biomedical Engineering, Texas A&M University, College Station, Texas 77843, USA

*Corresponding author: apple@tamu.edu

Received October 14, 2009; accepted November 17, 2009;posted January 5, 2010 (Doc. ID 118396); published February 10, 2010

We have developed a high-resolution molecular imaging technique, pump–probe optical coherence micros-copy (PPOCM), based on the fusion of pump–probe spectroscopy and optical coherence microscopy. We havedemonstrated the prototype system on a fixed human skin sample containing a nodular melanoma. The re-sults indicate that the PPOCM can clearly provide a strong contrast between the melanotic and amelanoticregions. Potential applications of the PPOCM imaging of melanin include the early diagnosis of melanomaand the mapping of tumor margins during excision. The technique may in general be applied to any biologi-cal chromophore with a known absorption spectrum. © 2010 Optical Society of America

OCIS codes: 170.3880, 180.4315, 170.1790.

High-resolution optical molecular imaging currentlyplays an essential role as a research tool for biology,biochemistry, and the biomedical sciences. The pre-vailing high-resolution optical molecular imagingmodalities based on fluorescence need exogenous tagsin order to probe the majority of biomolecular speciesbecause of their poor intrinsic fluorescence. In addi-tion to the increased experimental complexity, thesetags may potentially interfere with the process understudy or even prove toxic to the sample. The need forexogenous tags also impedes applications in humanssince any tag must garner FDA approval. It is there-fore desirable to develop high-resolution molecularimaging techniques that do not rely on fluorescenttags but rather exploit physical phenomena whichare more common in biomolecules [1,2].

Pump–probe spectroscopy is a well-established toolin molecular physics for measuring the spectrum anddynamics of molecular species that are poor fluoro-phores. The foundation of the technique lies in thedetection of transient changes in the probe attenua-tion induced by the pump radiation. While there area number of different schemes involving molecularexcited states, some of which we have enumeratedearlier [3], we are primarily interested in the tran-sient bleaching of the ground state induced and de-tected by degenerate or near degenerate pump andprobe wavelengths. In this manifestation of pump–probe spectroscopy the pump drives molecules to theexcited state thereby depleting the ground statepopulation. The intensity of the probe radiation isthen stronger when the pump is on as compared towhen the pump is off due to the reduced probabilityof absorption and the increased probability of stimu-lated emission.

We have recently adapted degenerate pump–probespectroscopy to optical coherence tomography (OCT)to yield ground state recovery pump–probe optical co-herence tomography (gsrPPOCT) [4]. We were able toimage a number of endogenous chromophores includ-ing hemoglobin in the efferent filament arteries of anadult zebrafish. We were also able to measure anduse the ground state recovery time of hemoglobin and

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tions from both in a mixture. The spatial resolutionsin the gsrPPOCT and OCT, 5–10 �m axial and10–20 �m lateral, are well suited to imaging themicrometer-scale morphology of living tissues.

The work described here represents our effortto extend the resolution of this technique beyondthe micrometer-scale morphology to the cellular/subcellular level. To realize that goal we have com-bined pump–probe spectroscopy with optical coher-ence microscopy (OCM) to yield pump–probe opticalcoherence microscopy (PPOCM). OCM is the fusion oflow coherence interferometry with confocal micros-copy. In the OCM, the scattered light is filtered byboth the coherence and confocal gates to yield scat-tering based images at depths superior to standardconfocal microscopy [5]. The addition of pump–probespectroscopy will introduce a molecular dependentgate, which will enable the simultaneous acquisitionof a molecular image and a scattering based image.

Image formation in PPOCM follows from the threegates enumerated above. The coherence and confocalgates combine to form the point spread function(PSF) of a typical OCM system. The coherence gatecontributes only in the axial dimension �z� and isgiven by PSFL�z�=exp�−z24 ln��z�−2�, where �z is theFWHM of the PSF and we have assumed a Gaussiansource spectrum. Note that �z is the coherencelength of the source, which is related to the sourcecenter wavelength and bandwidth by �z=2 ln 2�0

2�����−1, where �0 is the center wavelengthand �� is the source FWHM. The confocal gate con-tributes in both the axial and radial dimensions. Theconfocal PSF, following Sheppard and Gu [6] is givenby PSFC�r ,z�=I�r ,z�2, where I�r ,z� is the intensity inthe focal region. The PSF for OCM is then the prod-uct of the two,

PSFOCM�r,z� = exp�−z24 ln 2

��z�2 �I�z,r�2. �1�

In the two limiting cases, long coherence length orlong depth of focus, Eq. (1) becomes the PSF for a

confocal microscope or an OCT system, respectively.

2010 Optical Society of America

February 15, 2010 / Vol. 35, No. 4 / OPTICS LETTERS 533

The molecular gate in the PPOCM is due to theoverlap of the pump and probe beams. FollowingDong et al. [7], in the absence of a confocal gate thePSF due to the overlap of the pump and probe isgiven by PSFP-P=I�r ,z�I��r� ,z��, where the prime in-dicates the pump beam and the transition of thechromophores is assumed to be far from saturation.When combined with the confocal gate the PSF is im-proved to PSFP-P=I�r ,z�I��r� ,z��I�r ,z�. The additionof the coherence gate leads us to an equation for thePSF in PPOCM,

PSFPPOCM�r,z� = exp�−z24 ln 2

��z�2 �I�r,z�2I��r�,z��. �2�

Comparing Eqs. (1) and (2) it is clear that the spa-tial resolution of the PPOCM may exceed that of thebase OCM system depending on the focusing of thepump beam. In the limit that the pump beam is un-focused, PSFPPOCM is equivalent to PSFOCM. Simi-larly as we approach saturation of the transition,PSFPPOCM will approach PSFOCM. Even if the transi-tion is completely saturated by the pump along theentire beam path the spatial resolution will never de-teriorate below that of the base OCM system. This isimportant, since the maximum PPOCM signal isachieved when the transition is saturated. If ourtechnique relied solely on the overlap �PSFP-P�, themaximum signal would be accompanied by a drasticdeterioration of the spatial resolution.

A probe bandwidth larger than the bandwidth ofthe transition being probed will not improve thePPOCM resolution, because the additional band-width will not be modulated by the pump. Addition-ally, subresolving the confocal gate by decreasing thesource coherence length leads to an ambiguity in thespatial position where the absorption occurs, becauseabsorption is an integrative process in the pathlength. Recovering the depth dependent absorptionwould require the computation of the derivative im-age [4].

We have selected melanin as the initial target chro-mophore for our PPOCM system. Melanin is one ofthe most abundant chromophores in the human body.The different forms of melanin play roles in the func-tions of the brain, ear, eyes, and skin. There are anumber of potential applications for the imaging ofmelanin. One of the most important is as a tool to in-vestigate and potentially diagnose melanoma, a par-ticularly aggressive cancer starting in the melano-cytes and the most common form of skin cancer.

The optical system is shown schematically in Fig.1. A tunable bimodal source was generated bylaunching approximately 250 mW of light from a tun-able (680–1080 nm) 140 fs Ti:sapphire laser (80 MHz)into a 6.5 cm length of the polarization-maintaininghigh-NA single-mode fiber. The center of the spec-trally broadened pulse [Fig. 2(A)] was removed by fil-tering with a polarizer to yield the bimodal spectrumshown in Fig. 2(B). The probe light ��z�50 �m� waslaunched into an interferometer based on a Mach–Zehnder design. A piezoelectric transducer (PZT) in

the reference arm, driven at resonance (57.25 kHz)

by a sine wave, provided a carrier frequency for theinterferometric signal. The chopped (3.4 kHz) pumpbeam passed through an optical delay line, whichalso rotated the polarization by 90°. The pump andprobe, combined in a polarizing beam splitter, werefocused onto the sample using a 0.8 NA water immer-sion objective which provided a sub-0.5-�m lateralresolution and a measured �2 �m axial (confocalgate) resolution. A 2�2 (50/50) single-mode fiber cou-pler served as the confocal pinhole as well as the finalbeam splitter in the interferometer. The two outputsof the fiber coupler were connected to a 125 kHz dualbalanced detector. En-face imaging was accomplishedby translating the sample using a high precisionstage. The measured signal-to-noise ratio of the OCMsystem was 122 dB, 5 dB from the shot-noise limit of127 dB.

The intereferometric signal �SOCM� was extractedas in [8] by taking the square root of the sum of thepower at the PZT fundamental modulation frequency�P1� and the second harmonic �P2�, SOCM=�P1+P2. Aslong as the amplitude of the PZT motion was equal to0.84 of a fringe then SOCM as calculated above was in-dependent of the phase of the interference signal.The amplitude modulation of the pump beam intro-duced sidebands on P1 and P2 at +/− the pump modu-lation frequency when there was a pump inducedchange in the probe signal amplitude. The sidebands,corresponding to the PPOCM signal, were processedsimilar to the OCM signal, SPPOCM=�P1−+P2−

+�P1++P2+, where the + and � subscripts indicatethe plus and minus sidebands, respectively.

The sample, shown in Fig. 3(A), was prepared bycutting a �1 mm thick section from fixed human skincontaining a nodular melanoma perpendicular to thetissue surface. A region of tissue in the boundary area

Fig. 1. Schematic diagram of the pump–probe optical co-herence microscope: PMF, high-NA polarization-maintaining single-mode fiber; Obj, objective; BS, beamsplitter; PBS, polarized beam splitter; LPF, long pass filter;SPF, short pass filter; DM, dichroic mirror; 2�2, 2�250/50 fiber coupler.

Fig. 2. (A) Representative source spectrum after exitingthe PMF. (B) Representative source spectrum after the po-larizer. The short wavelength band was used as the pump,

and the long wavelength band was used as the probe.

534 OPTICS LETTERS / Vol. 35, No. 4 / February 15, 2010

between melanotic (left) and amelanotic (right) re-gions, indicated in Fig. 3(A), was scanned with thePPOCM system over a field of view of 135 �m�90 �m at a depth of �26 �m. The pump and probepowers on the sample were 2.16 mW and 1.17 mW,respectively, with the pump leading the probe by 50.5ps.

The resulting OCM and PPOCM images acquiredsimultaneously are shown in Figs. 3(B) and 3(C), re-spectively. The melanotic regions should have stron-ger attenuation due to absorption; however, in gen-eral they are also more highly scattering than theamelanotic regions because of the melanin granules.The net result is a fairly poor contrast between themelanotic and amelanotic regions in the OCM imagewith the melanotic regions appearing brighter. ThePPOCM image shows no signal in the amelanotic re-gions and a strong signal in the melanotic regions,thus providing a strong contrast. While the melaninis largely isolated on the left side, the PPOCM imageclearly indicates small “islands” of melanin depositsin the otherwise amelanotic region on the right.These islands are not discernable in the OCM image;however, they are consistent with the light micro-scope image in Fig. 3(A), which indeed shows smallblack areas infiltrating the amelanotic region on thesurface of the sample. The PPOCM image shows astrong molecular contrast and enables the identifica-tion of molecular deposits not clearly seen in theOCM image.

In addition to just providing molecular contrast, wewould like to map the relative concentration of thetarget chromophore. Unfortunately, while thePPOCM signal is a function of the chromophore con-centration �C� it is also a function of the tissue reflec-tivity �R�, pump power �Ppu�, and probe power �Ppr�.The PPOCM image is therefore a poor approximationof the relative chromophore concentration. At a fixed

Fig. 3. (Color online) (A) Light microscope image of thefixed nodular melanoma sample. The scale bar is 100 �m.The box indicates the approximate region where the OCMand PPOCM images were recorded. (B) OCM image. Thescale bar is 20 �m. (C) PPOCM image. (D) PPOCM/OCMratio image depicting the relative melanin concentration.Images (B)–(D) have the same field of view and scale.

depth, Ppu and Ppr reaching that depth is approxi-mately constant over small lateral distances; hencethe most important variable is R. We can take advan-tage of the fact that we also effectively measure thesample reflectivity simultaneously in the OCM im-age. The OCM image is formally a function of Ppr andR; hence the ratio PPOCM/OCM is a function of Cand Ppu and is a reasonable approximation of therelative chromophore concentration at a given depthin the tissue. This ratio is shown in Fig. 3(D). Beforetaking the ratio, both the OCM and PPOCM imageswere normalized and spatially averaged over 3 pixels�1.5 �m� in both dimensions to reduce the noise. ThePPOCM signal was additionally thresholded at 7% ofthe maximum intensity. While the spatial averagingwas insufficient to remove the speckle pattern fromthe OCM and PPOCM images, it is suppressed in theratio image, because the speckle is perfectly corre-lated in the OCM and PPOCM images. The PPOCM/OCM ratio image provides the relative chromophoreconcentration over small lateral distances and can-cels the speckle noise to yield a high-contrast low-noise molecular image.

In conclusion, we have developed a PPOCM systemand used it to image melanin in human skin tissuewith a nodular melanoma. The results indicate thatPPOCM can clearly provide a strong contrast be-tween the melanotic and amelanotic regions, which isnot available in the scattering based OCM image. Po-tential applications of the PPOCM imaging of mela-nin include the early diagnosis of melanoma and themapping of tumor margins during excision. While wehave demonstrated the prototype PPOCM system onmelanin the technique is more general and may beapplied to any biological chromophore with a knownabsorption spectrum and sufficient concentration.

We thank Michael Cohen, M.D., for supplying thefixed nodular melanoma sample. We gratefully ac-knowledge the financial support of this work throughthe National Institutes of Health (NIH)1R21RR025799.

References

1. D. Fu, T. Ye, T. E. Matthews, B. J. Chen, G. Yurtserver,and W. S. Warren, Opt. Lett. 32, 2641 (2007).

2. W. Min, S. J. Lu, S. S. Chong, R. Roy, G. R. Holtom,and X. S. Xie, Nature 461, 1105 (2009).

3. B. E. Applegate, C. Yang, and J. A. Izatt, Opt. Express13, 8146 (2005).

4. B. E. Applegate and J. A. Izatt, Opt. Express 14, 9142(2006).

5. A. L. Clark, A. Gillenwater, R. Alizadeh-Naderi, A. K.El-Naggar, and R. Richards-Kortum, J. Biomed. Opt. 9,1271 (2004).

6. C. J. R. Sheppard and M. Gu, Optik (Stuttgart) 86, 104(1990).

7. C. Y. Dong, P. T. C. So, C. Buehler, and E. Gratton,Optik (Stuttgart) 106, 7 (1997).

8. B. M. Hoeling, A. D. Fernandez, R. C. Haskell, E.Huang, W. R. Myers, D. C. Petersen, S. E. Ungersma,R. Wang, M. E. Williams, and S. E. Fraser, Opt.Express 6, 136 (2000).