Coherent Raman scanning fiber endoscopy
Brian G. Saar,1,4Richard S. Johnston,2Christian W. Freudiger,1,3X. Sunney Xie,1,5and Eric J. Seibel2,6
1Department of Chemistry and Chemical Biology, Harvard University, Cambridge, Massachusetts 02138, USA
2Human Photonics Laboratory, University of Washington, Seattle, Washington 98195, USA
3Department of Physics, Harvard University, Cambridge, Massachusetts 02138, USA
4Present address: MIT Lincoln Laboratory, Lexington, Massachusetts 02420, USA
Received March 28, 2011; revised May 7, 2011; accepted May 13, 2011;
posted May 13, 2011 (Doc. ID 144659); published June 20, 2011
Coherent Raman scattering methods allow for label-free imaging of tissue with chemical contrast and high spatial
and temporal resolution. However, their imaging depth in scattering tissue is limited to less than 1mm, requiring the
development of endoscopes to obtain images deep inside the body. Here, we describe a coherent Raman endoscope
that provides stimulated Raman scattering images at seven frames per second using a miniaturized fiber scanner, a
custom-designed objective lens, and an optimized scheme for collection of scattered light from the tissue. We
characterize the system and demonstrate chemical selectivity in mouse tissue images.
OCIS codes: 170.2150, 190.4180, 290.5910.
© 2011 Optical Society
Optical imaging methods are vital in medical diagnostics
because they offer high spatial resolution and excellent
contrast at high acquisition speeds compared to widely
used radiological imaging techniques such as magnetic
resonance imaging, position emission tomography, or
x-ray computed tomography. Optical disease diagnosis
can be enhanced in sensitivity by the application of
exogenous labels to obtain molecularly specific informa-
tion from the tissue, but these are sometimes toxic. In
addition, most biological tissues exhibit strong absorp-
tion and scattering of light, so the penetration of high-
resolution optical methods is limited to a few millimeters,
restricting optical diagnostics.
Endoscopy circumvents the penetration depth limit by
bringing an optical system into proximity with tissues of
interest, even inside a living patient. Optical endoscopes
with white light contrast have revolutionized surgery,
allowing minimally invasive procedures to be performed
have been reported for wide-field fluorescence and
narrowband imaging [2,3], confocal reflectance  and
fluorescence , and two-photon fluorescence [3,6–8].
The combination of novel label-free optical imaging
modalities with endoscopic devices presents new oppor-
tunities for in situ diagnostics. Coherent Raman scatter-
ing (CRS) [9–11] methods offer label-free chemical
contrast with microscopic resolution and have the poten-
tial to provide in situ virtual pathology on fresh tissue
without the excision, fixation, slicing, and staining proce-
dure required by traditional histopathology [12,13]. How-
ever, coherent Raman methods have required laser
scanning microscopes. Such systems can be miniaturized
, but the use of optical fiber delivery is highly desir-
able for flexibility in clinical applications. Several groups
have reported optical modeling  and progress toward
realizing optical-fiber-delivered CRS imaging systems,
but required either sample scanning  or scanning
mirrors  and long scan times. Here, we report a
miniaturized coherent Raman imaging system based on
a compact scanning fiber endoscope (SFE) , a step
toward translating coherent Raman imaging methods
into clinical practice. In CRS, which refers to both
coherent anti-Stokes Raman scattering (CARS) [9,10]
and stimulated Raman scattering (SRS) [11,19–22], laser
beams at two frequencies, called the pump, ωp, and
Stokes, ωS, frequencies, are used to illuminate the sam-
ple. When the difference frequency between the two
beams is tuned to match an intrinsic molecular vibra-
tional frequency in the sample, ωvib, several nonlinear
interactions occur. New light is generated at the anti-
Stokes frequency, ωas¼ 2ωp− ωS, by the CARS process,
and some intensity is transferred from the pump to the
Stokes beam by the SRS process [Fig. 1(a)]. Both pro-
cesses offer chemical selectivity, high spatial resolution,
and video-rate imaging speeds in vivo in living animals
 and humans . SRS offers more desirable contrast
that depends linearly on the target species concentration
and is free from the image artifacts and nonresonant
background that plague CARS imaging , but has thus
far not been implemented in an endoscope or miniature
Because SRS imaging involves measurement of small
intensity changes in the pump beam, we make use of a
highly sensitive modulation transfer scheme in which we
amplitude modulate the Stokes beam at a high reference
frequency (e.g., 10MHz), illuminate the sample with both
the pump and Stokes beams, and then detect only the
pump beam with a photodiode (PD) behind an optical
filter (OF) that blocks the Stokes beam. The PD signal
is demodulated by a lock-in amplifier (LIA) tuned to
the reference frequency that only detects amplitude
modulation when SRS occurs in the sample .
Images are formed by scanning the laser over the sam-
ple and recording the modulation of the pump beam as a
function of position. In this Letter, the laser system that
excites the sample is similar to one described recently
. It is based on a passively mode-locked Nd:YVO4
laser (picoTrain, high Q laser), which provides 7ps
pulses at a repetition rate of 80MHz. A portion of the out-
put at 1064nm is the Stokes beam for the CRS processes,
and the remainder of the output is frequency doubled and
used to synchronously pump an optical parametric
2396 OPTICS LETTERS / Vol. 36, No. 13 / July 1, 2011
0146-9592/11/132396-03$15.00/0© 2011 Optical Society of America
oscillator (OPO; Levante Emerald, APE GmbH), which
generates a tunable pump beam near 800nm.
The Stokes beam is modulated by an electro-optic
modulator (EOM; Raicol Crystals, Ltd.) at 10MHz and
combined in space on a dichroic mirror (DM; 1064dcrb,
Chroma Technology) with the OPO output. The two
beams are linearly polarized and, after passing through
an achromatic wave plate (AHWP05M-980, Thorlabs),
they are coupled into a single-mode polarization-
maintaining (PM) optical fiber (PM780-HP, Thorlabs).
The optical fiber delivers the laser beams to the sample
[Fig. 1(b)], and the image is scanned using a miniature
piezo-electric (PZ) actuator to drive the fiber tip at its
mechanical resonance frequency in an expanding spiral
[Fig. 1(c)]. A scan rate of seven frames per second can be
achieved. The scan pattern is actuated by passing the
optical fiber through a small (<0:5mm diameter) piezo-
electrically active ceramic tube that is radially patterned
with four electrodes. To achieve the spiral scan, one pair
of opposing electrodes is driven with a sine wave and the
other is driven with a cosine wave, both with linearly
expanding envelope functions. The scan driver electro-
nics track the fiber tip for image reconstruction. This
SFE provides a unique combination of small package
size, high frame rate, and adaptability to a variety of laser
The laser light from the fiber output is focused onto the
sample by a gradient index (GRIN) objective lens (OL).
Single element GRIN lenses have large axial chromatic
aberrations, which significantly degrade the nonlinear
signal-generation efficiency. For this reason, we utilized
an OL consisting of two GRIN elements sandwiching a
diffractive optic, which was designed to have zero axial
chromatic aberration at the chosen wavelengths, an NA
of 0.3, a free working distance of 150μm, and an outer
diameter of 1:4mm. Typical power levels of ∼130mW
are used, which is similar to those previously used for
in vivo imaging .
For light collection, we observed that the modulated
pump beam that travels forward from the focus is redir-
ected back toward the tissue surface by multiple scat-
tering within turbid tissue. Nonsequential ray-tracing
simulations [Fig. 1(d) and ] show that using a large
area PD surrounding the excitation lens allows for
optimal collection efficiency. In this study, we utilized
a 10mm × 10mm PD, but simulations indicate that a
5mm diameter circular PD would achieve ∼75% of the
present collection efficiency. Thus, we surrounded the
lens with a back-biased silicon PIN PD (FDS1010,
Thorlabs) through which we drilled a 2mm diameter hole
for the OL [Fig. 1(c)]. A custom-designed OF identical to
one in a previous report  blocked the modulated
Stokes beam from reaching the PD. As observed
previously , because picosecond pulses are employed
in CRS, ordinary silica fibers can be used for light deliv-
ery over a length of 1m without self-phase modulation
broadening the optical spectrum [Fig. 2(a)] beyond the
intrinsic Raman linewidth (∼1nm at 800nm).
We verified the spatial resolution of the system with
transmission mode CARS images of 2μm diameter beads
[Fig. 2(b)]. Bead sizes of 1:3μm (FWHM) demonstrate the
ability to clearly resolve 2μm objects across the 80μm
diameter field of view. By axially translating the beads
in 1μm steps, we measured the axial FWHM of 6:5μm.
We then tested the ability of the system to obtain SRS
images in mouse skin tissue at seven frames per second
(with a 500 pixel diameter field of view). This required
the use of a custom LIA  to process the SRS signal.
fiber input (black) and propagating through 1m of PMF with
the indicated power level at the fiber output. Even with
240mW of power output, the FWHM is 0:85nm, indicating that
we can still resolve individual Raman bands with intrinsic line-
widths of around ∼1nm. (b) CARS image of 2μm polystyrene
beads coated on a glass slide. Individual beads can be easily
resolved, indicating a lateral resolution better than 2μm. Scale,
10μm. (c) Measured lateral profile (red) and Gaussian fit (blue)
of an individual 2μm bead, indicating the lateral FWHM of
1:3μm for the bead indicated in green in (b). (d) Axial profile
of an individual 2μm bead, indicating an axial FWHM of 6:5μm
and demonstrating optical sectioning capability.
(a) Optical spectrum of the pump laser beam at the
(b) Schematic of the device: the Stokes laser beam is amplitude
modulated using an EOM driven by a 10MHz waveform,
temporally delayed, and combined with the pump beam on a
DM. The beams pass through a Glan-Laser (GL) prism, and their
polarization is rotated by a λ=2 plate to match the PM axis of a
1m PM fiber (PMF). Light is coupled into the fiber through an
OL (10× 0.4 NA Olympus). The distal tip of the fiber is moved
using a SFE. The scan is driven by control electronics through
electrical cables, which also report the scanner position to a
personal computer to record images. The photocurrent from
the detector in the SFE is demodulated using an LIA, which also
provides the reference frequency to the EOM. The demodulated
signal is the image intensity. For transmitted CARS detection,
the SFE tip is positioned in front of a photomultiplier tube de-
tector with an OF (labeled F) that transmits only the anti-Stokes
wavelength. (c) Detailed view of the SFE. The single-mode PMF
passes through a PZ tube patterned with four electrodes, which
are actuated by the scan controller. The light from the PMF
is relayed to the sample by a GRIN objective lens (OL) and
collected by a PD after filtering by an OF. (d) Simulations show
that, for tissue with scattering mean free path of 0:1–0:25mm,
forward-traveling SRS light is not efficiently collected by the
objective lens. Collection by a PD surrounding the detector
can achieve >30% efficiency of the light that propagates from
(a) Energy diagrams of the SRS and CARS processes.
July 1, 2011 / Vol. 36, No. 13 / OPTICS LETTERS2397
With this system, we could tune into the CH2stretching
vibration of lipids [Fig. 3(a)] to record an image of the
sebaceous gland [Fig. 3(b)] and subcutaneous fat in
mouse skin at a depth of ∼50μm [Fig. 3(c)]. Image con-
trast disappeared and the signal level dropped to the de-
modulator noise floor when the OL was not focused in
the tissue sample, indicating a negligible background
from nonlinear effects within the fiber.
We tested the chemical selectivity of the system by
recording images of mouse hair near the surface of
the tissue. Hair is known to be rich in protein and coated
with a layer of lipids that are secreted by the mouse skin.
By tuning the laser system to match the CH3stretching
resonance at 2950cm−1[Figs. 3(d) and 3(f)] or the CH2
stretching resonance at 2845cm−1[Figs. 3(e) and 3(g)],
we could see in the SRS images either the protein-rich
center of the hair [Fig. 3(d) indicated by the “H” labels)
or the lipid-rich coating of the hair [Fig. 3(e)]. This tuning
was previously observed in a bulk SRS microscope 
and demonstrates that the chemical contrast in tissue
matches the behavior of the Raman spectrum [Fig. 3(a)].
However, simultaneous forward CARS detection
[Figs. 3(f) and 3(g)] did not manifest this behavior be-
cause the nonresonant background of CARS limits the
chemical selectivity of the technique and does not allow
clear discrimination between lipids and proteins based
on CH2and CH3stretching images. This demonstrates
the importance of obtaining SRS contrast because recent
work utilizing this contrast has shown that SRS has
pathology-like diagnostic capability in fresh tissue .
In summary, the system described here demonstrates
that the technical challenges for CRS endoscopy, fiber
delivery, laser scanning, focusing, light collection, and
demodulation can be solved. Improvements to the
device, particularly the use of fiber lasers, will increase
ease of use, facilitating clinical translation of CRS.
We thank C. Lee and D. Melville for design and
fabrication of the SFE at the University of Washington;
B. Messerschmidt of GrinTech GmbH for design and
fabrication of the GRIN lens; and M. Roeffaers,
G. Holtom, D. Fu, and X. Zhang for helpful discussions.
C. W. Freudiger acknowledges support from a Boehrin-
ger Ingelheim Fonds Ph.D. scholarship. This work was
supported by National Institutes of Health (NIH) grants
to E. J. Seibel (R33 CA 094303 and R21 EB012666)
and a T-R01 grant (1R01EB010244-01) to X. S. Xie, the
Royalty Fund, and the Harvard University Office of
Technology Development Accelerator Fund.
1. M. J. Mack, J. Am. Med. Assoc. 285, 568 (2001).
2. K. Gono, IEEE J. Quantum Electron. 14, 62 (2008).
3. C. M. Lee, C. J. Engelbrecht, T. D. Soper, F. Helmchen, and
E. J. Seibel, J. Biophoton. 3, 385 (2010).
4. C. Boudoux, S. Yun, W. Oh, W. White, N. Iftimia, M.
Shishkov, B. Bouma, and G. Tearney, Opt. Express 13,
5. R. Kiesslich, J. Burg, M. Vieth, J. Gnaendiger, M. Enders,
P. Delaney, A. Polglase, W. McLaren, D. Janell, and S.
Thomas, Gastroenterology 127, 706 (2004).
6. M. E. Llewellyn, R. P. J. Barretto, S. L. Delp, and M. J.
Schnitzer, Nature 454, 784 (2008).
7. C. L. Hoy, N. J. Durr, P. Chen, W. Piyawattanametha, H. Ra,
O. Solgaard, and A. Ben-Yakar, Opt. Express 16, 9996
8. P. Kim, M. Puoris’haag, D. Côté, C. P. Lin, and S. H. Yun,
J. Biomed. Opt. 13, 010501 (2008).
9. A. Zumbusch, G. R. Holtom, and X. S. Xie, Phys. Rev. Lett.
82, 4142 (1999).
10. C. L. Evans and X. S. Xie, Annu. Rev. Anal. Chem. 1,
11. C. W. Freudiger, W. Min, B. G. Saar, S. Lu, G. R. Holtom,
C. He, J. C. Tsai, J. X. Kang, and X. S. Xie, Science 322,
12. C. L. Evans, X. Xu, S. Kesari, X. S. Xie, S. T. C. Wong, and
G. S. Young, Opt. Express 15, 12076 (2007).
13. C. W. Freudiger, B. G. Saar, D. A. Orringer, R. Pfannl,
Q. Zeng, L. Ottoboni, T. Chen, W. Ying, R. D. Folkerth,
C. A. French, W. R. Welch, C. Waeber, J. R. Sims, P. L.
De Jager, O. Sagher, M. A. Philbert, X. Xu, S. Kesari,
X. S. Xie, and G. S. Young, “Stain-free histopathology with
stimulated Raman scattering microscopy,” submitted to
Proc. Natl. Acad. Sci. USA.
14. S. Murugkar, B. Smith, P. Srivastava, A. Moica, M. Naji,
C. Brideau, P. K. Stys, and H. Anis, Opt. Express 18,
15. I. Veilleux, M. Doucet, P. Coté, S. Verreault, M. Fortin,
P. Paradis, S. Leclair, R. S. Da Costa, B. C. Wilson, and
E. J. Seibel, Proc. SPIE 7558, 75580D (2010).
16. F. Légaré, C. L. Evans, F. Ganikhanov, and X. S. Xie, Opt.
Express 14, 4427 (2006).
17. M. Balu, G. Liu, Z. Chen, B. J. Tromberg, and E. O. Potma,
Opt. Express 18, 2380 (2010).
18. E. J. Seibel, C. M. Brown, J. A. Dominitz, and M. B. Kimmey,
Gastrointest. Endosc. Clin. N. Am. 18, 467 (2008).
19. E. Ploetz, S. Laimgruber, S. Berner, W. Zinth, and P. Gilch,
Appl. Phys. B 87, 389 (2007).
20. Y. Ozeki, F. Dake, S. Kajiyama, K. Fukui, and K. Itoh, Opt.
Express 17, 3651 (2009).
21. P. Nandakumar, A. Kovalev, and A. Volkmer, New J. Phys.
11, 033026 (2009).
22. B. G. Saar, C. W. Freudiger, J. Reichman, C. M. Stanley,
G. R. Holtom, and X. S. Xie, Science 330, 1368 (2010).
23. C. L. Evans, E. O. Potma, M. Puoris’haag, D. Côté, C. P. Lin,
and X. S. Xie, Proc. Natl. Acad. Sci. USA 102, 16807 (2005).
mouse skin (black), a representative lipid compound (oleic
acid, red), and a representative protein compound (soy protein
extract, dotted magenta). (b) SRS image of a sebaceous gland
approximately 30μm deep in mouse skin. (c) SRS image of the
subcutaneous fat approximately 50μm deep in mouse skin,
acquired at seven frames per second in the epi direction. (d),
(e) SRS images of hairs near the surface of a mouse skin with
protein (d) and lipid (e) contrast. Hair is rich in protein and
coated in lipids, as indicated by the SRS images. (f), (g) CARS
images obtained simultaneously with (d) and (e) in the trans-
mitted direction, showing a lack of CARS contrast for protein
(f) and lipid (g). Field of view, 80μm; scale, 10μm.
CRS images with the endoscope. (a) Raman spectra of
2398 OPTICS LETTERS / Vol. 36, No. 13 / July 1, 2011