We present a single-beam coherent Raman microscopy method based on pump–probe, time-resolved stimulated Raman scattering (SRS) measurements with shaped probe pulses. In the single-beam method, we divide a broadband laser spectrum into three frequency bands for the pump, phase-modulated (PM) probe, and local oscillator (LO) probe pulses. Multiple low-wavenumber Raman modes are efficiently excited by an impulsive pump pulse, and a specific Raman mode can be selectively probed using temporal beam coupling between the PM and LO probe pulses via the Raman-induced refractive index modulation. To achieve both high sensitivity and a high spectral resolution, we allocate a large spectral bandwidth () to two probe bands and use a new selective detection scheme based on the spectral focusing technique. By giving a strong group delay dispersion to the probes (), we can obtain an improved spectral resolution of down to . In a proof-of-concept experiment, the intrinsic molecular-vibration contrast of sevoflurane, an inhaled anesthetic drug, is successfully visualized. This result suggests that single-beam SRS imaging with pulse shaping is a potentially powerful tool for detecting the Raman signals of small-molecule drugs in living cells and tissues.
© 2017 Optical Society of America
Optical microscopy provides us with rich physical and chemical information concerning living cells and tissues at sub-micrometer spatial resolutions. In particular, fluorescence microscopy has served as an invaluable tool in biomedical science due to its high detection sensitivity, high affinity, and specificity for molecular targets of interest. Despite these advantages, the use of fluorescent labels for targeting small-mass molecules (less than 300 Da) is often problematic. This is because most fluorescent labels are bulky, and the labels may alter the biological activity of the targeted small molecules . Another approach is to use a label-free vibrational contrast, such as Raman scattering  or infrared absorption , which allows us to record molecular-specific spectral signatures. For observing intact living organisms, Raman scattering is preferred because it is hardly affected by water absorption and provides sub-cellular spatial resolution. A key challenge to standard Raman imaging is how to increase the speed, that is, how to obtain more signal photons than background photons, because spontaneous Raman scattering is generally much weaker than fluorescence.
Coherent Raman scattering microscopy [4–6], an emerging imaging technology using nonlinear optical effects, such as coherent anti-Stokes Raman scattering (CARS) [7,8] or stimulated Raman scattering (SRS) [9–11], has increasingly gained attention due to its enhanced sensitivity. By enhancing the Raman signal via coherent amplification with ultrafast laser pulses, the sensitivity can be significantly improved by up to several orders of magnitude. CARS and SRS imaging have been successfully applied to various studies, such as lipid biology , transdermal pharmacokinetics [13,14], and non-labeled histopathology [15,16].
One practical issue for coherent Raman microscopy is related to its complicated laser system. The coherent Raman process occurs through optical excitation via the beating of two incident pulses at a specific Raman-resonant frequency , where and are the angular frequencies of the pump and Stokes pulses, respectively. To implement this system, therefore, one has to prepare a perfectly synchronized, two- or multi-wavelength, widely tunable pulsed laser source, such as an optical parametric oscillator , or an additional frequency broadening via an optical fiber [18,19]. This complexity has made it difficult for non-physics researchers to take advantage of coherent Raman technology.
An alternative, simpler approach is single-beam multiplexed CARS microscopy [20–25]. Instead of aiming at a specific Raman frequency with two-color narrowband pulses, one can excite and detect multiple Raman modes simultaneously using a single broadband laser source. In addition to improved simplicity and robustness, impulsive excitation has proven particularly efficient in the low-wavenumber spectral region due to pump–Stokes degeneracy, where the Raman modes exist within the broadband laser bandwidth. With this approach, we have developed a single-beam phase-modulated CARS microscope  and have successfully demonstrated a direct label-free detection of small molecules (the Raman peak of taurine at approximately ) inside a mouse cornea tissue using time-resolved heterodyne detection . However, the detection sensitivity of previous single-beam CARS microscopy has been limited compared to that of multi-beam CARS or SRS systems.
Theoretically, the signal-to-noise ratio (SNR) of coherent Raman microscopy in the shot-noise-limited regime is proportional to the pump power times the root square of the probe power () [6,11], and it reaches its highest value when the pump-to-probe power ratio is set to 2:1. In most single-beam CARS setups, however, the power of the probe beam is much smaller than desired. This is because there is a trade-off between the spectral resolution and the sensitivity. The spectral resolution of single-beam CARS is limited by the effective bandwidth of the probe beam, which is determined by the bandwidth of the filter element used for pump–probe separation, such as a bandpass filter [20,21], a spectral phase modulation gate [22,23], or a notch filter . However, given a broadband mode-locked oscillator with a smoothed spectral power-density profile used as a single-beam source, directly reducing the probe bandwidth to obtain a high spectral resolution results in the reduced power of the probe. Consequently, the SNR available in a conventional single-beam CARS setup has been restricted by unbalanced pump-to-probe power ratios.
In this paper, we propose a single-beam phase-modulated stimulated Raman scattering (PM-SRS) microscopy with improved sensitivity. We overcame the above mentioned trade-off using a novel pulse-shaping technique, which we refer to as spectrally focused detection. Spectral focusing is known to be a method for generating a narrowband pulse from broadband, frequency-locked pulses. This strategy has been widely used in a variety of nonlinear optics applications, such as terahertz pulse generation , sum frequency generation , and coherent Raman scattering measurements [28,29]. Here, our basic idea is to use spectral focusing for selective spectroscopic detection rather than for narrowband nonlinear excitation. We prepare a pair of frequency bands for probe pulses: one is phase modulated, and the other is used as a strong local oscillator (LO) for optical heterodyne detection. While using efficient impulsive excitation and allocating a relatively large bandwidth to the probe pulses to improve the sensitivity, we can still obtain a spectral resolution smaller than the probe bandwidth with the help of spectral focusing via frequency-locked probe pulses.
Here, we present the system design and characterization of our first prototype of a single-beam PM-SRS microscope with spectral-focusing detection. We also demonstrate a non-labeled spectroscopic imaging of a lightweight anesthetic drug. The new single-beam scheme presented here is a potentially powerful and robust method for detecting the Raman signals of small drug molecules in living cells and tissues.
A. Phase-Modulated Stimulated Raman Scattering
Figure 1 shows the schematics of amplitude-modulated SRS (AM-SRS) and PM-SRS. An SRS signal is observed via constructive or destructive spectral interference between a coherent Raman field and an LO field, namely, stimulated Raman gain (SRG) and stimulated Raman loss (SRL) [Fig. 1(a)]. In commonly used two-color AM-SRS, the intensity of either the pump or Stokes beams is modulated [Fig. 1(b)]. The intensity-modulated and non-modulated beams are combined such that the two-color pulses are spatially and temporally overlapped, generating an interference fringe in the ultrafast time domain [Fig. 1(c)]. When the beating frequency between the two-color pulses, , matches the Raman resonant frequency of a sample, the incident pulses actively drive the molecular vibration. Here, the two pulses are used not only for excitation, but also for probing the vibration via self-heterodyne detection. Owing to this excitation-detection degeneracy, the relative phase of the pump and Stokes pulses is always cancelled out, and eventually, a fixed phase lag ( or on Raman resonance) is left between the induced coherent Raman field and the LO field. This explains why an AM-SRS output signal is synchronized with the AM input regardless of the relative phase between the pump and Stokes pulses.
In the single-beam PM-SRS scheme, however, we divide a broadband laser spectrum into three frequency bands for the pump, PM probe, and LO probe pulses [Fig. 1(d)]. Instead of modulating the intensity, we apply a periodic saw-wave PM input from to and, therefore, a virtually linear phase to the PM probe pulses [Fig. 1(e)]. Our single-beam PM-SRS is based on a pump–probe configuration [Fig. 1(f)]. The pump frequency band shown in Fig. 1(d) is used to generate a femtosecond pulse for impulsive excitation, exciting multiple molecular vibration modes simultaneously. From a classical perspective, each vibrational response to the impulsive excitation field can be expressed as the temporal modulation of the effective refractive index. When the pulse width is much shorter than the lifetime of the vibrational relaxation , the vibrational response is approximately given as
The coherent excitation is then followed by the PM probe () and the LO probe (), as illustrated in Fig. 1(f). From multiply excited Raman modes, a targeted Raman frequency of is selectively probed by a phase-to-intensity conversion process using two-beam coupling via Raman-induced refractive index modulation in the temporal domain. When the probe frequencies and are tuned such that , and the probe pulse width satisfies , the PM-SRS signal detected at the LO probe band is derived to be
The signal detection mechanism represented by Eq. (2) can be regarded as heterodyne coherent Stokes Raman scattering (CSRS). In fact, the PM-SRS presented here is described as a special case of the dissipative coherent Raman process where heterodyne CARS and CSRS occur simultaneously. The resultant energy exchange between the two probe pulses is essentially equivalent to SRG and SRL. A clear difference from conventional AM-SRS is that the relative phase between the coherent Raman and LO fields () is not cancelled out in the pump–probe configuration. The Raman gain and loss is therefore alternately observed depending on the relative phase [Fig. 1(d)]. As a result, the PM-SRS output signal is synchronized with a saw-wave PM input [Fig. 1(e)]. More detailed quantum descriptions of this heterodyne detection scheme have been discussed in Refs. [4,30,31].
The AC intensity signal modulated at the PM frequency can be extracted as an output via lock-in detection (or demodulation). This modulation transfer scheme significantly reduces low frequency noise and improves the detection sensitivity performance. Moreover, the PM-SRS system is robust against phase fluctuations from thermal drift or the vibrations of the optics because the AC modulus output (the R-output of a lock-in amplifier) is insensitive to the phase offset, . This is achieved only when we apply a saw-wave or a linear phase modulation; any nonlinear PM input (e.g., sinusoidal modulation with resonant drive electronics) would generate spurious harmonic components that would severely impair the signal stability.
B. Spectrally Focused Detection
In a previous single-beam heterodyne CARS spectroscopy, a portion of a Fourier-transform-limited pulse (the non-resonant emission or excitation field itself) has been used as an LO, and a narrowband probe has been used to resolve the multiply excited Raman modes [20–25]. Our alternative strategy is to recover the unbalanced pump–probe power ratio due to the use of a narrowband probe by allocating a wide bandwidth to the PM and LO probe bands and by manipulating the spectral phase. Here, we show that spectrally focused detection with shaped probe pulses provides highly sensitive Raman signal detection with an improved spectral resolution.
Suppose we measure a specimen with Raman spectrum , and the targeted molecule is excited by the pump field . The refractive index modulation induced by the excited molecular vibration can be described as6) using a shifted auto-correlation function as 5) and (7) indicate that the coherent Raman spectrum can be measured by scanning . In the single-beam setup, this can be readily implemented by changing the allocation of the LO band [i.e., a flexible spectral phase control with a spatial light modulator (SLM)]. The spectral resolution of the selective detection is determined by the frequency width of the spectral cross correlation . In the case of transform-limited probe pulses, for example, we can write the cross correlation as 8), we can see that the spectral resolution is simply given by the root mean square (RMS) of the probe bandwidths. In this case, as is common in conventional heterodyne CARS with a narrowband probe, the trade-off relationship between the probe power and the resolution cannot be avoided.
This trade-off can be circumvented by shaping the spectral phase of the two probes. As indicated in Eq. (5), the key requirement is that the frequency width of the probe cross correlation be smaller than that of the bandwidth-limited resolution. To obtain a sharp cross correlation of the probe pulses, we take the spectral focusing approach sketched in Fig. 2. Here, we assume the simplest spectral focusing scheme, where the same amount of linear chirp and group delay are applied to both probe bands. When the probe beat frequency matches the Raman resonant frequency (), the induced coherent Raman field is perfectly phase matched with the LO field [Fig. 2(a)]. Therefore, the spectral interference becomes in-phase over the entire LO band, and the probe cross correlation reaches its maximum [Fig. 2(b)]. When the probe beat frequency is detuned (), the frequency of the induced field is de-centered by , and a relative delay is generated between and [Fig. 2(c)]. This delay, , leads to a spectral interference fringe [Fig. 2(d)]. The net integrated SRS signal described in Eq. (5) then sharply decreases with the frequency mismatch. The frequency pitch of the spectral interference is given as5), or the cross correlation , reaches the first null when the fringe pitch approaches the probe bandwidth . Therefore, the spectral resolution of the spectrally focused detection is given by
From this expression, we find that the spectral resolution can be improved by providing a sufficiently large GDD to the probe pulses. For example, if a probe centered at 780 nm with a bandwidth of 10 nm (corresponding to ) is linearly chirped by a GDD of , we can obtain a fine spectral resolution of . This indicates that we can allocate a larger bandwidth without losing spectral resolution by using spectrally focused detection, enabling the efficient use of the energy budget of the single-beam setup.
C. Single-Beam PM-SRS Microscope
The single-beam PM-SRS microscope system consists of a single broadband Ti:sapphire laser (Vitara, Coherent: 80-MHz repetition rate, 15-fs pulse width, centered at 790 nm), a newly developed phase control unit and microscope, and monochromator optics for signal detection (Fig. 3).
The former part of the phase control unit is standard dispersion compensation optics that contain chirped mirrors for GDD compensation and a 4f pulse shaper with an SLM (LCOS-SLM, X10468-02, Hamamatsu: 800 pixel) for higher-order dispersion compensation (see Appendix A for details). Another important role of the SLM pulse shaper is arbitrary frequency allocation and pulse shaping of the LO band. The LO probe band (centered at between 790 nm and 860 nm, 10 nm bandwidth) is defined as the frequency region in which the spectral phases (delay, linear, and nonlinear chirp) are manipulated to perform the spectrally focused detection.
The latter part of the phase control unit is for high-speed phase modulation of the PM probe. The pulses in the PM probe band, which is defined as the passband of a bandpass filter [FBH780-10, Thorlabs: centered at 780 nm with a 10-nm bandwidth ()], are phase modulated by an electro-optic modulator (EOM) (EO-PM-NR-C1, Thorlabs). The EOM is driven by a high-voltage saw-wave input modulated at 35 kHz with a rise–fall duty of 95%, generating a virtually linear -phase modulation. The temporal delay of the PM probe is adjusted by moving the end mirror of the PM probe, whereas that of LO probe is adjusted by giving the equivalent first-order spectral phase on the SLM. The same delay value is provided to both the PM and LO probe pulses. The PM probe and the reflected pump pulses (including the LO probe) are re-combined at the same bandpass filter and focused onto a sample via a microscope objective lens (M Plan Apo NIR NA 0.50, Mitsutoyo).
After the transmitted pulses are collimated by a second identical objective lens, the LO frequency band for the SRS output is selectively extracted via 4f monochromator optics to increase the SNR. The detection bandwidth and the center frequency are tuned with a variable slit on a translational stage at the intermediate Fourier plane. The filtered pulses are detected using a silicon PD (Newfocus 2031, Newport). The PD output is demodulated by a lock-in amplifier (LIA) (LI 5640, NF Corporation), followed by simultaneous recording of the DC and the LIA outputs with a digitizer (USB-6009, National Instruments) that provides the transmission and the PM-SRS images, respectively. Laser scanning is performed by driving an XYZ-positioning sample stage. The stage scanning, signal acquisition, and SLM phase manipulation are controlled using the Labview platform (National Instruments).
A. Selective Spectroscopic Detection
First, we evaluated the spectral resolution of the PM-SRS microscopy with the spectrally focused detection. Pure dimethyl sulfoxide (DMSO) enclosed in a quartz cuvette was used as a test sample. The SRS spectrum was acquired by scanning the center frequency of the LO band , which was done by simply changing the SLM phase pattern. The detection bandwidth of the monochromator optics was set equal to the probe bandwidth (), and the center frequency was synchronously tuned with the SLM frequency scan to increase the SNR.
To achieve a high spectral resolution, we designed and manipulated the spectral phase of the PM and LO probe pulses. In the PM probe band, a temporal delay was added by directly changing the optical path length, and linear and nonlinear chirps were passively provided by the band-pass filter. Therefore, no SLM-based manipulation was added to the PM probe (see Appendix A). In the LO probe band, the delay, linear, and nonlinear chirps were fully controlled by the SLM pulse shaper. To reproduce the measured phase profile of the PM probe shown in Fig. 8(a), a GDD of and a third-order dispersion (TOD) of , which was derived by a polynomial fitting of the PM probe phase, were added to the LO probe. The delay of both the probes was set to .
Figure 4(a) plots the spectrum obtained with the spectrally focused detection. We can see a sharp Raman peak of DMSO at . The full-width at half-maximum (FWHM) of the Raman peak is estimated to be , which corresponds to a higher spectral resolution compared to the probe bandwidth. This result roughly agrees with the theoretical resolution expressed in Eq. (10) ( ( in FWHM) for and ).
Even though we successfully achieved an improved spectral resolution in terms of the FWHM, some residual non-resonant background (approximately 7% of the Raman peak intensity) was still visible in the off-resonant region. In addition, we observed an unwanted spectral sidelobe near the Raman peak. To reduce the background and the sidelobe, we provided a larger delay (1530 fs) to the probes and a larger GDD () and TOD () to the LO band to “defocus” the spectral selectivity. The detection bandwidth of the monochromator was reduced to because the frequency near the filter edges provided little contribution to the SRS signal in this case. Using this defocused setting, we obtained a smooth Raman spectrum profile with a spectral resolution of in FWHM [Fig. 4(b)]. We confirmed that using a non-shaped LO probe results in a reduced resolution close to bandwidth limited (). The spectrally focused detection in this case still provided an improved resolution by a factor of 2 compared to non-shaped probing. We also confirmed that the non-resonant background was suppressed down to 1.5% in the off-resonant region due to the larger probe delay.
While giving a larger delay could further increase the signal-to-background ratio  and improve the detectability, it would also reduce the absolute signal level [Eq. (2)] and degrade the imaging speed. In practice, therefore, it is important to optimize the probe delay depending on the application.
In designing the spectrally focused probes, note that the achievable spectral resolution and delay time of the LO probe are limited by the finite spatial bandwidth of the 4f pulse shaper (Appendix B). The cutoff delay time, at which the space–time coupling loss (or the vignette at the objective lens) reaches 50%, is 2.2 ps in the current setup. Higher spectral resolutions will be available when using a high-resolution SLM and 4f optics with a sufficient spatial bandwidth.
B. PM-SRS Signal versus Chemical Concentration
To quantify the detection sensitivity of PM-SRS, we measured the SRS signal intensity for different concentrations of DMSO in water. The exposure powers of the pump (including the LO band) and the PM probe pulses on the sample were 26 and 3.0 mW, respectively. The measured SRS signal level was proportional to the square of the total incident power (Appendix C). The pulse-shaping parameters used in this experiment were the same as those used in Fig. 4(b).
Figures 5(a) and 5(b) depict the calibration curves derived from the recorded SRS signal intensities at . An excellent linear dependence of the SRS signal on the concentration, which is the advantage of using a strong LO field, was successfully confirmed. Note that the error bars in Figs. 5(a) and 5(b) represent the residual non-resonant background intensity (measured between 500 and ) that causes stationary spectral interference with the resonant Raman signal. The temporal fluctuation (or RMS noise level) of the SRS output in this measurement was in the order of (with a 1-s integration time), which hardly affects the total amount of error. At 1.0% DMSO (140 mM), we obtained a signal-to-background ratio of 2.0. The corresponding PM-SRS signal and background levels in the AC/DC intensity ratio () were and , respectively. From the result, the limit of the detectability at which the signal-to-background ratio reaches unity is estimated to be 0.5% (70 mM).
C. Vibrational Imaging
To demonstrate the non-labeled imaging capability of the PM-SRS microscope, we imaged a transparent droplet of pure sevoflurane, an inhaled anesthetic drug, in water. In the vibrational imaging experiment, a small amount of liquid sevoflurane and water was injected and enclosed into a quartz cuvette, and the sample was placed on the laser focal plane. A two-dimensional near-infrared transmission image and an SRS image (image size: , field of view: ) were simultaneously captured using XY-raster stage scanning with a linear velocity of . A total of 2000 samples were sequentially acquired along the -direction with a sampling rate of 10 kHz and then reduced to 200 pixels by binning (pixel integration time: ), where the time constant of the LIA output was 0.1 ms. The number of lines along the -direction was 200 (0.5-μm steps).
Figure 6(a) depicts the vibrational spectrum of sevoflurane measured by PM-SRS, where the pulse shaper settings were similar to those used in Fig. 4(b) (delay: 1120 fs, GDD: , TOD: , detection bandwidth: ). The Raman resonance peak of sevoflurane was observed at approximately . The peak SRS signal level and water background level (primarily from modulated non-resonant background) were and , respectively.
Figure 6(b) shows a near-infrared transmission image (DC output). We observed a strong negative contrast around the droplet attributable to the difference in the refractive index between sevoflurane and water. Figure 6(c) shows a raw SRS output image captured at the resonant frequency (). Even though a clear vibrational contrast was obtained at this point, the negative contrast was still reflected in the raw SRS image as shown in Fig. 6(c). Nevertheless, in a ratio SRS image normalized by the DC output shown in Fig. 6(d), such a negative-contrast artifact was successfully removed. The SNR of the normalized SRS image (defined as the ratio between the averaged SRS signal in the sevoflurane droplet region and the standard deviation in the water background region) was calculated to be 6.2. The vibrational contrast disappeared in the off-resonant image (), as shown in Fig. 6(d), which clearly demonstrates the molecular-specific imaging capability of this modality.
Our PM-SRS setup has several advantages over previous single-beam CARS systems:
- (i) Spectrally focused detection eliminates the drawback of the poor pump-to-probe power ratio in conventional single-beam methods, providing a better spectral resolution without sacrificing sensitivity.
- (ii) Heterodyne SRS detection with a strong LO ensures a linear relationship between the chemical concentration and the signal level, which enables an analysis based on quantitative imaging.
- (iii) More importantly, heterodyne detection removes the stationary background, including the coherent non-resonant background (proportional to ) and incoherent backgrounds, such as two-photon-excited fluorescence. In homodyne CARS (with no LO), the shot noise of these background effects ultimately limits its detectability.
- (iv) Time-resolved pump–probe measurements allow the selective detection of a long-lived, narrowband Raman signal from drug molecules, while suppressing the short-lived Raman background from macromolecules in a tissue or cell, as well as the non-resonant background .
It is worth noting that recent investigations of a spectrally tailored single-beam CARS system (Brückner et al. [33,34]) have reported improved spectral resolution and sensitivity (i) and time-resolved detection (iv), although the quantification capability and detectability are still limited because it is based on homodyne CARS.
Another advantage of the PM-SRS imaging system is the compatibility with our previously reported single-beam CARS spectroscopy [20,21], which is also based on impulsive excitation and heterodyne detection. This compatibility is important because it allows us to work on imaging and spectroscopy studies on the same platform. In addition, the impulsive excitation setup can be upgraded to a multiplex SRS imaging system by simply replacing the monochromator with multiple bandpass filters.
Our single-beam PM-SRS microscope is aimed at detecting a small-molecule drug hidden behind the background noise in living cells or tissues. In a biological setup, the detectability of standard CARS and SRS is still limited to tens of mM or more . Detecting sub-mM concentrations is strongly desired, yet currently only achieved by imaging coupled with small Raman tags that have a resonant peak in a cell-silent spectral region . A time-resolved, pump–probe method seems to be the most promising strategy to suppress the Raman-background from macromolecules in a non-labeled manner.
The major limitation of the current prototype is the residual non-resonant background () whose amplitude is proportional to the temporal pump–probe overlap, . In the present study, we focused on impulsive excitation to achieve the highest excitation efficiency for low-wavenumber Raman modes. The modulated non-resonant background could be further reduced by combining the spectrally focused detection with other efficient excitation methods, such as multi-pulsed excitation by sinusoidal spectral-phase modulation  or phase-coded excitation . Targeting high-wavenumber modes () can also be achieved by using a sub-10-fs broadband laser [19,33,34]. Another technical task is to reduce the sidelobe level in the cross-correlation trace. One can find that the presented selective detection scheme based on cross correlation is a spectral analogue of the matched filtering technique, which has been used in a variety of fields, including radar systems  and optical image correlators . Further suppression of the background and sidelobe could be achieved using a more sophisticated matched-filter design.
Time-resolved SRS spectroscopy has been used to probe vibrational modes in the excited electronic states of pigments or fluorescent molecules , where the first pulse is used for electronic excitation and the second and third pulses are used for SRS measurements. In PM-SRS, by contrast, only the Raman modes excited by the pump pulse are probed, whereas self-induced Raman modes excited by the PM and LO probe pulses are rejected. In fact, the latter self-heterodyned SRS effect does occur in our PM setup. However, it simply results in a DC intensity offset and provides no observable AC signal. Therefore, PM-SRS probes purely time-resolved Raman signals, which have never been achieved by conventional AM-SRS. Even though delaying the probes reduces the overall sensitivity to some extent, the detectability (or the signal-to-background ratio) of drug molecules with a long relaxation lifetime (a few ps or more) could be substantially improved.
Another background mechanism that may limit detectability in standard SRS microscopy includes AM-induced backgrounds, such as cross-phase modulation, photo-thermal lensing effects, and transient absorption . In PM-SRS, however, such an AM-induced artifact is estimated to be negligibly small because the measured intensity modulation of the PM probe is in the order of .
In conclusion, we have developed a new single-beam coherent Raman microscope and characterized its performance, including its spectral selectivity, sensitivity, detectability, and vibrational imaging capability. The mechanism of the pump–probe SRS technique, which we refer to as PM-SRS, is understood to be a special case of the dissipative coherent Raman process, where non-self-heterodyned CARS and CSRS occur simultaneously.
The trade-off between sensitivity and spectral resolution, which has restricted the use of single-beam coherent Raman microscopy, has been overcome by spectrally focused detection with shaped probe pulses. We demonstrated that this new selective detection scheme provides a fine spectral resolution ( in our current setup), even when combined with efficient impulsive excitation. Like AM-SRS, the modulation transfer scheme using high-speed phase modulation provides high SNR and the heterodyne SRS detection ensures an excellent linear relationship between the signal and the chemical concentration. In a proof-of-concept imaging experiment, the molecular-specific vibrational contrast of a transparent anesthetic drug was successfully visualized. These results are supported by precise on-sample pulse measurement and a flexible spectral phase control technique that can be implemented with a standard SLM-based pulse shaper. We anticipate that the presented time-resolved SRS scheme could potentially improve the detectability of non-labeled coherent Raman microscopy.
APPENDIX A: PULSE MEASUREMENTS
Because single-beam coherent Raman microscopy uses a broadband pulse, precise spectral phase measurement and control are essential. In our single-beam PM-SRS setup, the spectral phase of the pump pulse is measured and pre-compensated using multi-photon intra-pulse interference phase scanning (MIIPS), an SLM-based pulse measurement method . The advantage of MIIPS over other measurement techniques is its on-sample measurement capability, which is particularly desirable in nonlinear microscopy applications.
In MIIPS measurements, a type-I BBO crystal is placed on the sample plate, and the sum-frequency generation (SFG) intensity is recorded with a spectrometer (USB2000, Ocean Optics). Careful calibrations of the SLM and the spectrometer were performed prior to the measurements. In addition, before a direct on-sample measurement, we confirmed that even a strong negative chirp was successfully compensated for by MIIPS [Fig. 7(a)]. The residual phase error after the MIIPS compensation relative to a commercial SPIDER system  (FC-SPIDER, APE) was 0.36 rad RMS between 750 and 840 nm [Fig. 7(b)], corresponding to a near-transform-limited pulse with a 19.5-fs pulse width. The relative phase error was presumably caused by imperfect phase wrappings on the SLM due to the limited spatial bandwidth of the 4f optics. In the actual on-sample MIIPS, we can avoid most of the phase wrappings by presetting the chirped mirrors such that the positive GDD of the intermediate optics, and the objective lens is compensated for right at the sample. Therefore, the accuracy of the on-sample MIIPS measurement could be better than the test result shown here. Figure 7(c) plots the SLM input profile that was obtained from the on-sample MIIPS. We can see that no phase wrapping occurred in the scanning range of the LO probe band ().
In fact, when we initially tried measuring the entire pump pulse using MIIPS, we observed a spurious multi-peaked SFG spectrum and could not obtain the converged output data even after several MIIPS iterations (not shown). From the data analysis, we found that this was caused by the bandpass filter for the pump–probe separation, which had steep phase jumps near the cutoff and cut-on wavelengths. To avoid this error, we separately measured the spectra of the longer wavelength above the cutoff and the shorter wavelength below the cut-on. After both MIIPS measurements (five iterations for each), we further adjusted the relative delay and the zeroth-order phase offset between the longer and shorter pump frequency bands so that the total SFG output reached the maximum. The SLM pulse shaping pattern in Fig. 7(c) was obtained as a result of these procedures.
The spectral phase of the PM probe was measured using a similar on-sample SFG setup; however, this time, we used an alternative algorithm (phase-resolved interferometric spectral modulation: PRISM ) that enables high-resolution spectral phase measurements. We confirmed that the PM probe indeed had a strong nonlinear chirp and a steep phase profile near the spectral edges at 775 and 785 nm [Fig. 8(a)], which can be explained by the dispersion characteristics of a Bragg reflector . The temporal pulse shape of the PM probe was also double-checked via SFG cross-correlation measurements, where the relative delay of the near-transform-limited pump and the PM probe was scanned by moving the end mirror of the probe arm. As predicted by the PRISM data, the recorded PM probe pulse had an asymmetric temporal profile [Fig. 8(b)]. The asymmetric pulse had a sharp rise and a long tail toward the delayed time direction. Fortunately, such an asymmetric pulse shape is preferred when probing the Raman vibrational coherence close to the pump pulse while minimizing the non-resonant background. For this reason, we apply no additional phase compensation to the PM probe band, whereas the phase profile of the LO probe band is adjusted to reproduce that of the PM probe.
APPENDIX B: BANDWIDTH OF THE 4F PULSE SHAPER
The available amount of temporal delay and chirp and, in turn, the maximum frequency resolution in a SLM-based 4f pulse shaper is limited by the spatial bandwidth, which is determined by two limiting factors.
The first limitation is related to the spatial frequency bandwidth of the SLM device and is defined as
The second limitation is the spatial bandwidth of the 4f optics, defined as3), and is the focal length of the 4f optics. In our current setup (, ), the spatial bandwidth is calculated to be .
The ratio of the bandwidths, , gives the spatial beam overlap between the shaped and non-shaped pulses after the pulse shaper, where half of one beam overlaps the other beam when . In the current setting, , which means that the spatial bandwidth limit is dominated by the 4f optics. This 4f-optics-limited bandwidth of corresponds to half of the full SLM resolution (4 pixels per cycle) and determines the spatial frequency cutoff, where the optical throughput decreases by half due to imperfect beam overlap (loss by vignette at the objective pupil). The cutoff spatial frequency is then converted to the cutoff delay time by the following time–space relationship:B3), the cutoff delay time of our current setup is derived to be , which agrees well with the measured cutoff delay time (2.2 ps).
APPENDIX C: POWER SCALING AND SNR LIMIT
According to Eq. (2), the SRS output signal is theoretically proportional to and therefore to the square of the total incident laser power. We validated the quadratic power scaling of the PM-SRS signal by changing the incident beam power with a neural density (ND) filter inserted prior to the microscope. We confirmed that the estimated scaling factor was nearly quadratic () [Fig. 9(a)]. We also verified the linear power scaling of the detection by changing the detected beam power with an ND filter inserted after the microscope, and the result showed a perfectly linear dependence [Fig. 9(b)]. This result confirms that the AC signal (the lock-in amplifier output) normalized by the DC level is proportional to the sample concentration, regardless of the amount of optical loss after the sample focal point.
Note that the detection sensitivity of PM-SRS is ultimately limited by the shot noise of the LO field, which is proportional to . Therefore, the theoretical SNR limit of PM-SRS in the shot-noise-limited regime scales as .
Japan Agency for Medical Research and Development (AMED).
This work was supported by the Japan Agency for Medical Research and Development. The authors thank T. Suzuki and M. Hayashi for the design and development of the instruments and H. Mogi for the helpful technical discussions. The authors are also grateful to M. Kawagishi and S. Terada for valuable suggestions for the experiment.
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