In fiber delivered SHG and TPEF microscopy, the spectral changes induced by fiber nonlinearities through SPM and SRS are often a primary concern. In CARS microscopy, which is based on the use of two pulse trains, fiber nonlinearities produce additional frequency components in addition to the spectral changes to the individual pulses. It has been shown that when pump (ωp
) and Stokes (ωS
) pulses are time-overlapped in optimized PCFs, four-wave-mixing processes can generate strong contributions at the anti-Stokes frequency (2ωp
] Depending on the phase-matching properties of the PCF, anti-Stokes components can be generated at substantial shifts up to 3000 cm−1
, a range of direct relevance to lipid imaging. To optimize the design of the fiber delivered probe, we first examined the effect of these fiber nonlinearities on the quality of the CARS excitation light.
Pulse delivery without significant spectral and/or temporal broadening is an important criterion for selecting a delivery fiber. Although it has been shown that pulse broadening effects are minimal in standard silica SMFs with lengths of less than 1m for picosecond pump and probe pulses with energies of a few nJ[14
], such fibers do not support femtosecond pulse trains with pulse energies relevant to CARS microscopy. To avoid spectral broadening effects, PCFs have been used as the primary excitation delivery fibers in multiphoton microscopy.[18
Here we examined a double-clad photonic crystal fiber (DCPCF16, Crystal Fibre) for use in multiphoton CARS microscopy. This fiber was selected because it could potentially be used for both efficient delivery of laser light and subsequent collection of back-scattered CARS signals. Double-clad PCFs offer the possibility of collecting the back-scattered signal through the fiber cladding, while the excitation beam is delivered through the fiber core, which makes these fibers ideal for general optical imaging endoscopy applications.[20
] This fiber has a 16 µm core that minimizes optical nonlinearities, which has enabled its use in endoscopic nonlinear optical imaging applications.[9
] As expected, limited spectral broadening of both the femtosecond and picosecond pulses was observed in this fiber of length less than 1m, as depicted in .
Intensity spectra of the pump beam measured before and after the (a) DCPCF16 fiber (FWHM=9.7 nm before the fiber; FWHM=9.7 nm after the fiber) and (b) the LMA-20 PCF (FWHM=10 nm before the fiber; FWHM=11 nm after the fiber)
The third fiber we tested was a large mode area PCF (LMA-20, Crystal Fibre) of 20 µm core diameter. This fiber has previously been used as a delivery and detection fiber in optical coherence tomography (OCT).[22
] In this fiber, no significant spectral pulse broadening was observed, as illustrated in . While the temporal duration of the picosecond Stokes beam was unaffected, the fs pump increased in duration by a factor of 2.5 (from 280 fs to 700 fs).
While the spectral broadening was not a concern for the individual pump and Stokes pulses in the PCF fibers, new frequency components may arise when the pump and Stokes are temporally overlapped in the fiber as a result of frequency mixing. In particular, for fibers that support phase-matching over a wide bandwidth, anti-Stokes frequency components can be generated through a FWM process in the fiber. To suppress such FWM effects, we selected PCF fibers that do not support phase-matching of the frequency components shifted by ~3000 cm−1
relative to the zero dispersion wavelength of the fiber. Hence, our fibers fulfill the condition:
are the wave vectors of the pump, Stokes and anti-Stokes components, respectively, and L
is the length of the fiber over which the frequency components remain temporally overlapped. Under these conditions, no coherent anti-Stokes generation is expected through a nonresonant FWM process. Despite the phase mismatch between the frequency components, we observed significant anti-Stokes generation in the PCF fibers. The spectral content of the anti-Stokes shifted radiation is shown in . The isolated anti-Stokes component shows a well-defined spectral profile that corresponds to the spectral convolution of the pump and Stokes pulse spectra. Because no additional broadening of this shifted component is observed, we conclude that this contribution is generated directly through a nonlinear mixing process between the pump and the Stokes pulses, and thus independent of the SPM mechanism. Importantly, we observed an identical anti-Stokes component in the case of the silica SMF, confirming that the anti-Stokes radiation is not the result of accidental phase-matching in the PCF fiber. We verified that the intensity of the anti-Stokes component scales quadratically with the pump light and linearly with the Stokes radiation, confirming that this shifted contribution is the result of a four-wave-mixing process.
Spectrally-resolved anti-Stokes four-wave-mixing signal measured at the output of (a) the LMA20 fiber (FWHM=7.9 nm) output and (b) the silica SMF (FWHM=8.6 nm).
FWM in fibers using two pump beams of different color is a well-known mechanism of generating new frequency components around the zero dispersion wavelength.[23
] However, such mechanisms typically rely on phase-matched conditions and the limited width of the Raman gain spectrum in silica, resulting in only moderate spectral shifts (≤ 500 cm−1
) relative to the input beams. Specially tailored PCFs with exceptionally broad phase-matching conditions have been used to achieve FWM generation of components shifted as much as 3000 cm−1
] Large shifts under non-phase-matching conditions can be achieved through cascaded stimulated Raman scattering, producing an array of periodically spaced spectral components.[25
] Such spectral patterns are not observed in our experiments, suggesting that stimulated Raman processes based on the fundamental Si-O modes are not the primary source of the observed anti-Stokes radiation. A possible explanation for the observed FWM component is the population of higher lying vibrational states of Si-O overtones and fiber impurities through stimulated Raman pumping, followed by incoherent anti-Stokes scattering by the pump. The incoherent anti-Stokes light is sustained in the fiber as it is not affected by phase-matching with the pump and Stokes beam.
In all fibers tested here, the anti-Stokes component generated in the fiber was much stronger than the typical CARS signals generated in biological samples. We compared the strength of the anti-Stokes FWM signal to the CARS signal generated in a DMSO sample after the fiber. The comparison between the two signals as a function of the time delay between the pump and the Stokes pulses is presented in for the double-clad PCF and for the LMA-20 PCF. A two meter long double-clad PCF was also used in an effort to separate the fiber anti-Stokes component from the CARS signal generated in the fiber due to a longer walk-off distance of the pump and Stokes propagating pulses. However, we found this measure to be ineffective: the FWM signal from the fiber could not be sufficiently suppressed without significantly affecting the CARS signal generated in the sample.
CARS signal intensity from the DMSO sample and FWM from the fiber as a function of time delay between the pump and the Stokes beam for the DCPCF16 fiber (a) and the LMA-20 fiber (b).
Given the observed fiber nonlineartities, we have chosen to use a photonic crystal fiber because of its favorable dispersion properties relative to a standard single mode optical fiber. The presence of intrinsic anti-Stokes generation in the fiber necessitates spectral filtering of the excitation light before focusing it into the sample. We have, therefore, chosen to implement separate fibers for delivery of the excitation light and collection of the signal.[11
] We selected the large area PCF (LMA20) for laser pulse delivery because of its excellent suppression of spectral broadening effects. The 4X/0.1 NA objective used for fiber coupling provided a coupling efficiency of 40% for 817 nm and 20% for 1064 nm. Since maximum efficiency of this fiber corresponds to 780 nm, a higher coupling efficiency for 817 nm than for 1064 nm was expected. The average power after the fiber was ~80 mW for both beams. The anti-Stokes radiation generated in the fiber was filtered out by a dichroic mirror placed after the fiber. The back-scattered forward generated CARS signal in the sample was collected by a second fiber. The collection fiber was chosen to be a large mode area, multi-mode fiber for maximum collection. A collection efficiency of 80% was obtained by matching the 0.39 NA of the fiber with the 0.4 NA of the objective used for fiber coupling.
To assess the performance of our fiber-delivered probe for CARS imaging, we imaged three different biological tissues ex vivo. We chose to take images of skin and eyelid, superficial tissues that would be easy to access in future in-vivo imaging. Subcutaneous fat (5(a)), individual adipocytes () and meibocytes () are clearly resolved with high contrast. The contrast observed is comparable to the contrast seen in CARS images obtained through free-space detection of the back-scattered light.[8
] This suggests that the detected signal includes the back-scattered, forward generated CARS radiation, and that the contrast is not dominated by aperture effects at the detection fiber. In addition, the contrast is not affected by spurious anti-Stokes components from the delivery fiber, resulting in images that originate solely from CARS generation in the tissue.
Fig. 5 CARS images of thick tissue samples ex vivo at 2842 cm−1.a) Small adipocytes of mouse ear skin. b) Adipocytes of subcutaneous layer of rabbit skin tissue. C) Meibomian gland in mouse eyelid. Images were acquired in 2s. Scale bar is 50 µm. (more ...)
The present images demonstrate that our design of separate delivery and detection fibers provides a simple yet efficient approach for acquiring high quality CARS images free of spectral artifacts and aperture effects. Although our present design produces acceptable CARS images, further miniaturization will be required to optimize its use as a hand-held probe suitable for clinical studies. Based on the current scheme, such optimization can be achieved by incorporating MEMS scanners and miniature lenses.[13