628 OPTICS LETTERS / Vol. 30, No. 6 / March 15, 2005

Photothermal image flow cytometry in vivo

Vladimir P. ZharovPhilips Classic Laser Laboratories, University of Arkansas for Medical Sciences, Little Rock, Arkansas 72205

Ekaterina I. Galanzha and Valery V. TuchinSaratov State University, Saratov 410012, Russia

Received June 7, 2004

The capability of photothermal (PT) microscopy to image moving, unlabeled cells in real time in vivo is dem-onstrated in a study of circulating red and white blood cells in blood and lymph microvessels of rat mesen-tery. Potential applications of this optical tool, called PT flow cytometry, are discussed. © 2005 Optical So-ciety of America

OCIS codes: 1170.0110, 170.0180, 170.1530, 170.6920.

The imaging of single cells in vivo is potentially im-portant for the early diagnosis of diseases (e.g., can-cer) or for the study of the influence of various factors(e.g., drugs, radiation) on individual cells. This ap-proach, however, has not yet been established.1 Thephotothermal (PT) technique has shown the capabil-ity to visualize nonfluorescent absorbing cellularstructures at the living, single-cell level in vitro with-out labeling,2,3 as well as to characterize liquid flowwithout markers and detect moving, absorbing het-erogeneities in capillaries.4 However, PT imaging(PTI) of moving cells, especially in vivo, has not yetbeen realized to our knowledge.

The scanning of a focused laser beam across a cell,as in confocal microscopy,1 is not suitable for obtain-ing PT images of moving cells because the process istoo time consuming (taking at least seconds if notminutes) because cells cross the area of detection in0.1–0.01 s for even the relatively slow flow in capil-laries. To develop PT flow cytometry (PTFC), we usedthe principle of nonscanning PTI2,3 with an advancedsetup5 modified by the incorporation of high-speedphotosensors, new imaging algorithms, and accesso-ries for in vivo studies (Fig. 1). Briefly, we realize PTIby irradiating a single cell (or group of cells) with apump laser pulse [optical parametric oscillator (OPO)with a wavelength of 420–570 nm, a pulse width of8 ns, and a pulse energy of 0.1–400 mJ from Lotis,Ltd.]. Laser-induced temperature-dependent varia-tions of the refractive index in the cell are detectedwith phase-contrast imaging techniques (OlympusBX51 microscope with a CCD camera; AE-260E, Apo-gee, Inc.) with a second, collinear probe laser pulse(Raman shifter with a wavelength of 639 nm, a pulsewidth of 13 ns, and a pulse energy of 2 nJ). The lowenergy level chosen for the probe laser pulse permitsobjects to minimize the influence of probe-beam ab-sorption and saturation of optical sensors. The diam-eters of the pump- and probe-beam spots, withsmooth intensity profiles controlled by the CCD cam-era, range from 20 to 50 mm and 15 to 50 mm, re-spectively, and thus cover entire objects. As a result,formation of a PT image requires only one pumppulse, and hence time-consuming scanning is unnec-essary. The PT image is calculated as the difference

between the two probe-pulse images: one before and

0146-9592/05/060628-3/$15.00 ©

a second after the pump pulse with a tunable timedelay.2,3

We have chosen rat mesentery from among variousanimal models (e.g., mouse ear or skin) because of itsunique anatomic structure consisting of thin, trans-parent, duplex connective tissue with a single layerof blood and lymph microvessels. Rats were narco-tized and then placed on a customized, heateds37.7 °Cd microscope stage, with constant diffusion ofwarm Ringer’s solution (37 °C, pH 7.4).

Transillumination digital microscopy (TDM) madeit possible to identify red blood cells (RBCs) travelingthrough blood vessels and lymphocytes in lymph flow(Fig. 2). With a video recording we found that in mostintact microlymphatic vessels (diameter of50–150 mm) lymphocytes in flow had an average ve-locity of ,211±11 mm/s s0–800 mm/sd within theirperiodic oscillations (forward and backward flow). Incomparison, the one-way velocity of RBCs was sig-nificantly higher, up to 2 mm/s in blood vessels witha diameter of 20–30 mm, although in capillaries witha diameter less than 10 mm the RBC flow velocityranged from 100 to 500 mm/s. The high spatial reso-lution of TDM (300 nm at 1003, N.A. of 1.25 with im-mersion) allowed us to estimate cell size and even theshape (Fig. 3, left). In addition, PTFC (navigated byTDM) allowed us to obtain images of moving lympho-cytes and RBCs (Fig. 3, right) that showed structuresspecific to PT images (invisible with TDM) and asso-

Fig. 1. Schematic of the integrated laser PT

microscope–spectrometer.

2005 Optical Society of America

March 15, 2005 / Vol. 30, No. 6 / OPTICS LETTERS 629

ciated with the spatial distribution of absorbing cel-lular chromophores (e.g., hemoglobin in RBCs or cy-tochromes in lymphocytes).2,3 In the case of RBCs thePT image depended slightly on changes in the cells’spatial orientation that were likely due to their spe-cific shape (Fig. 3, bottom left). PTI of the cells wasperformed at a pump–pulse energy that caused nonotable cell damage.3,6

To partly reduce the difficulties of PTI in vivo (e.g.,the influence of refractive and absorbing back-grounds from connective tissues and vessel walls, orchanges in cell position), several approaches were at-tempted: (1) use of broader laser beams to cover alarger area of vessels and ensure laser exposure ofsingle cells as they crossed the irradiated area, (2)study of small-diameter lymph and blood vesselswith relatively slow flow velocity, (3) use of nanosec-ond pump laser pulses synchronized with the actualcell velocity, (4) change in the position of the focalplane to make it possible to image cells traveling indifferent cross sections of lymph vessels, and (5) useof an algorithm to subtract individual probe and PTimages.

With regard to approach (3), laser pump pulseswere triggered by signals from a photodiode whenrelatively rare slow-moving cells in a microlymphaticvessel crossed the He–Ne laser beam and thuschanged its intensity (Fig. 1). When many cells trav-eled fast and close to one another (e.g., RBCs in bloodmicrovessels), we also used a pump–pulse repetitionmode with a pulse rate of 10 Hz. In combination withbroader laser beam diameters, even withoutsynchronization, this mode produced images of indi-vidual cells at a flow velocity of less than 0.5 mm/s,although it led to multiple exposures (and hence im-aging) of the same slow-moving cells.

With regard to approach (5), the first PT image wasobtained before the cell crossed the irradiated area;the subsequent PT image was obtained while the cellcrossed this area. This procedure made it possible toreduce the influence of absorption of the pump pulseby the vessel wall or connective tissue. It was foundthat laser-induced refractive heterogeneities around

Fig. 2. Bottom, typical image of rat mesentery with bloodand lymph microvessels. Top left, RBCs in a capillary. Topright, cells in lymph flow.

strongly absorbing objects (e.g., RBCs) exceeded the

natural refractive background from the vessel wall,as well as phase distortion of the probe beam itself.In other words, in the probe–pump–probe algorithmdescribed earlier, the PT image and the second probeimage (obtained 10 ns after the pump pulse) providedsimilar high-contrast images. This approach made itpossible to quickly obtain images of strongly absorb-ing objects by use of only one probe pulse to qualita-tively identify them among low-absorbing objects.

With its relatively slow frame rate, TDM could notdistinguish individual RBCs in relatively high-speedblood flow (Fig. 4, first, top row); however, PTFC withan 8-ns pump pulse allowed observation in the sameblood vessels of some laser-induced thermal effectsaround individual RBCs among cell clusters withoverlapping thermal fields (Fig. 4, second, top row).The PT images of vessels without cells (achieved bysqueezing the vessels) revealed no significant, local-ized heterogeneities (Fig. 4, first, bottom row), al-though in some cases faint structures were observed.In these cases we tried to find another position for

Fig. 3. Optical transmission (left column) and PT (rightcolumn) images in vivo of a single, moving lymphocyte (toprow) and RBC (bottom row) in lymph flow in rat mesentery(vessel diameter, 105 mm, velocity, ,120 mm/s). Pump–pulse parameters: wavelength, 525 nm; pulse width, 8 ns;pulse energies, 30 and 0.5 mJ (top and bottom, respec-tively); time delay, 10 ns.

Fig. 4. Top row, optical (first) and PT (second) images ofblood microvessels with a relatively high flow velocitys.1 mm/sd. Bottom row: first, PT image of blood vesselswithout cells; second, PT image of microlymphatics with20-nm gold nanoparticles. The dashed circles show the po-sition of the probe beam. Pump laser parameters: wave-length, 525 nm; pulse width, 8 ns; pulse energy, 0.5 mJ(second, top row) and 100 mJ (first and second, bottom

row); time delay, 10 ns.

630 OPTICS LETTERS / Vol. 30, No. 6 / March 15, 2005

the pump beam on vessels so that the PT image didnot include such structures.

We also demonstrated the capability of PTFC to vi-sualize 20-nm gold nanoparticles injected into alymph microvessel (Fig. 4, second, bottom row) as amodel for in vivo testing of the concept of selectivekilling of cancer cells by laser-induced microbubbleformation around gold nanoparticles attached tocells.6 Single nanoparticles appeared as diffraction-limited spots on PT images because bubbles ex-panded beyond the nanoparticles’ size to the diffrac-tion limit,6 although PT images from someaggregated particle clusters or large bubbles ex-ceeded the diffraction spots.

Potential applications of PTFC in vivo may include(1) identification of cells with significant differencesin natural absorptive properties (e.g., the counting ofwhite cells in blood flow or of rare RBCs in microlym-phatic vessels under normal and pathologic condi-tions); (2) monitoring of the circulated cells labeledwith PT nanoprobes for both diagnostic and thera-peutic purposes6; (3) study of laser-cell interactions,especially heat-diffusion dynamics in flow for thepurpose of optimizing laser treatment of vascular le-sions; and (4) study of the influence of different fac-tors (e.g., drug, glucose, and radiation) on blood cellsas demonstrated in vitro.3 Currently, the PTFC rateof ,10 cells/s is limited by the repetition rate of thepump laser s10 Hzd. Nevertheless, compared with theexisting nonflow PTI mode (which requires at least5 s to obtain one PT image), PTFC is approximately50 times faster. This capacity is enough to detect ev-ery nucleate white cell circulating in 20-mm bloodvessels at a typical rate of 3–6 cells/s (for a flow ve-locity of 2 mm/s), as well as all the cells in microlym-phatic vessels in which flow velocity is slower. Thedetection efficiency (detected cells per total numberof cells passing the vessels) in small lymphatics andblood capillaries with a laser beam diameter compa-rable with that of the vessels is expected to be close to1, whereas for vessels with a larger diameter there isstill a problem with obtaining clear images of each in-dividual cell. The ultimate goal of detecting, withoutconventional labeling, circulating metastatic cancercells at a desired threshold of 1:106 requires a signifi-cant improvement in the PTFC speed of analysis (un-der the current conditions that allow only slow flowto be studied, one cancer cell may appear in an irra-diated area only once a day) and additional evidencethat PTI can reliably discriminate cancer cells fromnormal cells. Nevertheless, a model for detecting can-cer cells labeled with PT probes (e.g., by their injec-tion into the localized primary tumor) at a relatively

high threshold s1:104d may be useful and realistic.PTFC may also be realized in vitro under more con-

trollable conditions than those for in vivo studies.The approach of combining the PT technique withflow cytometry (FC)7 as an alternative or supplemen-tal method to scattering and fluorescent techniqueshas not yet been analyzed. Indeed, despite its manysuccessful applications, conventional FC cannot cap-ture images of moving cells, which is crucial to mor-phologic analysis. The typical PT response has an ex-ponential temporal shape with a decay time of5–10 ms depending on the cooling time of the cell asa whole.6 Thus, theoretically, the maximum speed ofPTFC may be 105 cells/s, which is comparable withhigh-speed FC with hydrodynamic flow focusing.7

The realization of high-speed PTFC will require theuse of already routinely used lasers (e.g., argon, dye,or diode lasers) in a cw mode with 10–40 W of powerand an electro-optic (or other) modulator that pro-vides a laser pulse duration of 0.1–1 ms with arequired single-pulse energy level of 1–10 mJand a pulse rate of 0.1 MHz. In addition, severalcommercially available cameras now provide104 and even more images per second (e.g., Cooke orNAC Corporation), although it is sufficient to achievea more realistic and economic speed of 103 frames/s(i.e., 103 cells/s) in combination with an ellipsoidalbeam cross-sectional profile.7 In general, PTI can beapplied to the imaging of other relatively transparentstructures in vivo (e.g., C. elegans; this work is cur-rently underway in our laboratory).

This work was accomplished at the University ofArkansas for Medical Sciences and supported byNational Institutes of Health grant EB001858 andin part by National Science Foundation grantBES-0119470 to V. P. Zharov (zharovvladimirp@uams.edu). The authors thank Dmitri O. Lapotkofor his significant contribution to the development ofthe PT microscope.References

1. W. Rudolph and M. Kempe, J. Mod. Opt. 44, 1617(1997).

2. D. Lapotko, G. Kuchinsky, M. Potapnev, and D.Pechkovsky, Cytometry 24, 198 (1996).

3. D. Lapotko, T. Romanovskaya, and V. Zharov, J.Biomed. Opt. 7, 425 (2002).

4. V. P. Zharov, in Laser Analytical Spectrochemistry, V. S.Letokhov, ed. (Hilger, Boston, 1986), p. 229.

5. V. Zharov and D. Lapotko, Rev. Sci. Instrum. 74, 785(2003).

6. V. P. Zharov, V. Galitovsky, and M. Viegas, Appl. Phys.Lett. 83, 4897 (2003).

7. A. L. Givan, Methods Cell Biol. 63, 19 (2001).