Flexible Minimally Invasive CARS Measurement Method with Tapered Optical Fiber Probe for Single-Cell Application

We proposed and demonstrated a exible, endoscopic, and minimally invasive coherent anti-Raman Stokes scattering (CARS) measurement method for single-cell application, employing a tapered optical ber probe. A few-mode ber (FMF), whose generated four-wave mixing band is out of CARS signals, was selected to fabricate tapered optical ber probes, deliver CARS excitation pulses, and collect CARS signals. The adiabatic tapered ber probe with a diameter of 11.61 µm can focus CARS excitation lights without mismatch at the focal point. The measurements for proof-of-concept were made with methanol, ethanol, cyclohexane, and acetone injected into simulated cells. The experimental results show that the tapered optical ber probe can detect carbon-hydrogen (C–H) bond-rich substances and their concentration. To our best knowledge, this optical ber probe provides the minimum size among probes for detecting CARS signals. These results pave the way for minimally invasive live-cell detection in the future.


Introduction
The ability to detect intracellular biomolecules and track changes in their concentrations is essential for revealing many prevalent human diseases. For instance, lipids in foam cells play a signi cant role in determining plaque vulnerability of atherosclerosis 1 , and increased de novo lipid synthesis in the cells is a prevalent feature of human malignancies due to the dysregulated metabolic pathways 2 . Detecting the presence and concentration of biomolecules in living cells is essential for studying disease progression and early diagnosis. Coherent anti-Stokes Raman scattering (CARS) is a technique suitable for detecting intracellular biomolecules because of its advantages of chemical selectivity with label-free detection, which results from probing targeted molecular vibrational information [3][4][5] . CARS is an ideal tool for detecting carbon-hydrogen (C-H) bond-rich biomolecules such as lipids, which show distinct vibrational features in the 2700 cm −1~3 200 cm −1 of CARS spectrum.
Conventional CARS systems use high numerical aperture (NA) objective lenses to converge the pump light and the Stokes light onto the samples. Such systems achieve the research on highly arti cialisolating individual cells and the quantitative characterization of intracellular biomolecules [6][7][8][9] . Although studying single cells is helpful, cells in biological tissues also need to be studied much because it is well known that cells behave quite differently in biological tissues than they do when separated in arti cial environments 10 . However, due to the inhomogeneous refractive index (RI) of biological tissues, focused excitation lights are highly scattered and thus cannot reach the cells deep into biological tissues. Moreover, conventional CARS systems require perfect collimation, which is challenging to maintain.
The optical ber as a laser transmission carrier overcomes some of these restrictions, which can deliver the excitation lights into the sample directly and closely 11 . Owing to the exibility and portability, the optical ber is feasible for CARS endoscopic optical detecting cells deep inside tissues. To reduce the sizes of CARS ber-optic endoscopic probes, micro-scanners and small lens systems are applied to substitute high NA objective lens for focusing excitation lights [12][13][14][15][16] . Although these approaches have been demonstrated in endoscopically detecting the CARS signals deep in the tissue, the millimeter size of these endoscopes prevents them from enabling the non-tissue-destructive, in vivo chemical analysis tools 17 . Advances in optical design, including the use of metamaterials, novel ber designs, and 3Dprinted or freeform optics, enable further miniaturization of CARS optical ber endoscopic probes 10 .
Lombardini et al. inserted a silica microsphere into the output end face of an optical ber by using a CO 2 laser splicer to fabricate a compact CARS optical ber probe. The silica microsphere served as a lens and focused two output excitation lights on samples 18 . However, the silica microsphere focused the pump light and Stokes light at different positions due to the inherent chromatic aberration, resulting in reduced CARS excitation e ciency. Meanwhile, the sphere shape is di cult to puncture cell membrane and realize intracellular detection, and the process of xing glass microspheres on the ber facet is challenging.
The tapered optical ber probe is a promising solution for these problems. First, the tapered optical ber probe can be easily fabricated using a homemade or commercial tapering machine. Second, the tapered optical ber probes will not cause immense damage to cells due to their small size, and the insertion, interrogation and removal processes are time-saving. Furthermore, extracellular events such as cell secretion can be easily detected due to the exibility of tapered optical ber probes 19 . The tapered optical ber probes are widely employed in bioanalysis and cellular-related applications, such as intracellular PH detection 19,20 , cellular protein detection 21,22 and cancer metabolic compound detection 23 . In addition, the light eld with a large intensity gradient emitted by the tapered optical ber probes can be used for optical ber tweezers 24 . In other words, the CARS signals detection and optical capture of cell organelles can be integrated into a single tapered optical ber probe.
In this article, we proposed and demonstrated a exible endoscopic minimally invasive CARS measurement method with a tapered optical ber probe, as shown in Fig. 1(a). A few-mode ber (FMF) with generated four-wave mixing (FWM) band out of CARS signals was selected for adiabatic taper optical ber probe fabrication and ber deliver line. The adiabatic taper optical ber probe was used to realize CARS excitation lights focusing without mismatch at the focal point. We theoretically analyzed the fraction of the optical power in the ber, as well as the dispersion and nonlinearity of the optical ber probe, which indicates that fabricated tapered optical ber probe with a distal cladding diameter of 11.61 µm enable to deliver and focus the excitation lights, and cause no signi cant time domain broadening or spectral shift on the excitation pulses. The measurements for proof-of-concept were made with methanol, ethanol, cyclohexane, and acetone injected into simulated cells. The experimental results show that the tapered optical ber probe can detect C-H bond-rich substances and their concentration. To our best knowledge, the optical ber probe provides the minimum size among probes for detecting CARS signals.
The proposed tapered optical ber probe is promising to be a non-tissue-destructive, in vivo chemical analysis tool, and opens a signi cant route towards minimally invasive CARS signals detection of living cells deep in tissues.

CARS excitation and signals collection systems
Our homemade CARS excitation source can generate two chirped ultra-short pulse trains to be utilized as the pump light and the Stokes light, and the CARS excitation source is shown in Fig. 2 The CARS signals collection part collects and detects the generated forward-CARS (F-CARS) signals of samples, including an FMF as collection ber, a collimator, a short-pass lter (cutoff wavelength 750 nm, FESH0750, Thorlabs), and a photomultiplier tube (PMT, PMM02, Thorlabs). The F-CARS signals of samples are collected by the collection ber, ltered out from the excitation pulses by the lter, and detected by the PMT.

CARS tapered optical ber probe design
The CARS tapered optical ber probe was fabricated with an FMF. Since the mode eld diameter of the FMF is larger than that of the single-mode ber (SMF), the nonlinear effects in the ber core of the FMF are weaker. Compared with the step-index FMF and the multimode ber (MMF), the graded-index FMF supports fewer modes and effectively reduces the impact of intermodal dispersion on CARS excitation lights. Thus, we fabricated a tapered optical ber probe using a graded-index FMF (FM GI-4, YOFC). The core and diameter of the FMF are 23 µm and 125µm, respectively. Moreover, the utilization of the gradedindex FMF ensures coupling e ciencies of 86% and 78% for the pump light and the Stokes light, respectively. In the case of SMF, the coupling e ciencies are ~70% and ~60% for the pump light and the Stokes light, respectively.
In addition, the FMF-based tapered ber probe system can realize CARS excitation delivering and CARS signals detection without background FWM signals. The spectral region from 2700 cm −1 to 3200 cm −1 represents the 'C-H window' in Raman spectroscopy. In this region, various biological intracellular molecules such as lipids and proteins possess unique characteristics. However, tapered optical ber probes and ber delivery systems for CARS detection may encounter the problem of generating FWM signals in the ber itself at the same spectral region, which cannot be discerned from the useful CARS signals. The FWM signals of a common graded-index MMF are shown as the red solid line in Fig. 3. The FWM signals of the MMF cover from ~2700 cm −1 to 2900 cm −1 and interfere with the detection of C−H bond vibrations. Thus, the ber probe and ber delivery systems using such ber need to suppress or eliminate the generated FWM signals for detection without background noise. Although micro dichroic mirrors, lters and dual-wavelength waveplates can be used to suppress or eliminate FWM signals 13,16,26,27 , these spatial optics elements would make the ber probes bulky, which are incompatible with deep tissue detection in vivo. FMF probes can achieve noise-free CARS detection without ltering elements and reduce the size of probes for the reason that the FWM background of the FMF we used covers from ~3400 cm −1 to 3800 cm −1 , as shown in Fig. 3. A possible explanation for the FWM signals of the FMF in the particular band (~3400 cm −1 to 3800 cm −1 ) is the impurities doped in the ber. The impurities are added into the core for RI adjustment, and the resonance peaks of some unique impurities are in the range of 3400 cm −1 to 3800 cm −1 , such as boron oxide 28 . Different conditions of ber processing and preparation may also cause the resonance peaks of the impurities to appear in the range of 3400 cm −1 to 3800 cm −1 . For example, commonly doped impurities within the optical ber, such as germanium dioxide, can be prepared in various methods. The resonance peak of the hydrothermallygrown germanium dioxide is near 3500 wavenumbers 29 .
The optical ber can be processed into non-adiabatic taper and adiabatic taper. Compared with the nonadiabatic taper, the adiabatic taper we adopted in this work has a smooth taper-transition region, resulting in the main part of the optical energy remaining in the fundamental mode without transferring to higher-order modes, avoiding the effects of intermodal dispersion on the CARS excitation process. As the ber is tapered down, the light can no longer be con ned in the core and is then guided by the claddingexternal medium boundary when the difference between the refractive indices (RIs) of the core and the cladding is not large enough. As the ber is tapered down further, the light can no longer be con ned in the ber and is leaked into the external medium if the diameter of the taper distal is too small, leading to increased insertion loss and reduced CARS excitation e ciency. To choose proper distal cladding diameters, we calculated the fraction η of the optical power con ned in the ber with different ber cladding diameters by using the nite-difference time-domain (FDTD) method. The cross-sectional diameter of the FMF cladding is in the range of 4 µm~125 µm. Since the taper length of the adiabatic ber taper is much larger than the ber radius, the taper diameter can be considered to decrease approximately linearly, and the ratio of cladding diameter to core diameter in the taper area remains constant at 23:125. The RI of human tissues is used as the RI of the external medium in the calculation. However, the RIs of human tissues are complicated, and the different tissues possess different RIs. Most human soft tissues have a higher RI than water (1.33), in the range of 1.35~1.45 30 . The RIs of water and human adipose tissues (ATs) are used as the RIs of the external mediums, representing the lower and upper limits of the RI of most human soft tissue. Because the RI of ATs is relatively high in the human body. As with any other materials, the RI of human tissue is strongly dependent on the light wavelength.
According to the literature 31 , the RI of the human ATs at a body temperature of 37°C was estimated to be 1.455 at 780 nm and 1.452 at 1030 nm.
The calculated η is shown in Fig. 4. The blue solid and dashed lines represent the η at 780 nm and 1030 nm when the RI of water is used as the RI of the external medium. The red solid and dashed lines represent the η at 780 nm and 1030 nm when the RI of ATs is used as the RI of the external medium. When the external medium is water, more than 90% of the energy can be con ned in the optical ber with 4 µm~125 µm cladding diameters. Whereas when the external medium is ATs, more than 20% of the light energy is leaked into the external medium if the cladding diameter is less than 11 µm, resulting in a loss of light energy and a reduction in the excitation e ciency of the CARS signal. When the external medium is Ats, and the ber diameter is 12 µm, the η is 92.4% at 780 nm and 85.3% at 1030 nm. Considering the fraction of power in the ber and the strength of the ber probe, a tapered optical ber probe with a ~12 µm diameter distal is suitable for the CARS tapered optical ber probe.

CARS tapered optical ber probe fabrication
The tapered optical ber probe was processed using the hydrogen/oxygen ame brushing technique. The tapering process is: rstly, x the FMF to the xture and strip the coating layer from the area to be tapered. Secondly, the FMF is continuously heated and stretched by the ame. Thirdly, a ber splicing system (LDS 2.5, 3SAE) is used to nd the thinnest point of tapered ber, measure the taper waist diameter, and slice it into two parts. The taper length is 7 mm, and the taper diameter is 11.61 µm. The manufactured CARS tapered optical ber probe has a long pigtail (~50 cm) to quickly link with the optical power coupling device and facilitate operation. The fabricated tapered optical ber probe is shown in Fig. 1(b).

Calculation of the effect of tapered optic ber probes on CARS excitation pulses
To analyze and estimate the time domain broadening and spectral shift caused by the tapered optical ber, we calculated the dispersion and nonlinearity of the FMF with 4 µm~125 µm cross-sectional cladding diameters, and the results are shown in Fig. 5(a). The red solid line and the dashed red line represent the dispersion at 780 nm and 1030 nm, respectively. As a function of cladding diameter, the dispersion at 780 nm follows a similar trend to that at 1030 nm. Firstly, the dispersion at 780 nm and 1030 nm steadily remains -103.91 ps/nm·km and -25.73 ps/nm·km when the cladding diameter is larger than 30 µm, respectively. Secondly, the dispersion gradually decreases as the cladding diameter decreases. The dispersion at 780 nm reaches a minimum of -124.04 ps/nm·km with a cladding diameter of 18 µm and at 1030nm reaches a minimum of -41.44 ps/nm·km with a cladding diameter of 21 µm. Thirdly, the dispersion increases as the cladding diameter further decreases due to the increased proportion of light energy transmitted in the cladding and the larger diameter of the optical ber mode eld. In Fig. 5 W −1 /km at 1030 nm when the cladding diameter is 12 µm. Thirdly, as the cladding diameter further decreases, the nonlinearity turns to increase. As the cladding diameter decreases, the nonlinearity rst increases, then decreases, and nally increases again because the propagation mode transfers from the core mode to the cladding. The light spreads and redistributes in the cladding in the process, leading to a change of the effective mode area and ultimately to a change in the nonlinearity.
The general nonlinear Schrödinger equation (GNLSE) was used to numerically analyze the variation of the time domain broadening and the spectral shift of the ultrashort pulses through the tapered optical ber probe. The ber length is 7mm. We use the calculated dispersion and nonlinearity with 11.61 µm~125 µm cladding diameter in Fig. 5 to approximate the ber's characteristics at arbitrary positions along the ber. The calculated results are shown in Fig. 6. The red solid and dashed lines represent the time domain broadening at 780 nm and 1030 nm, respectively. According to the calculated results, the time domain broadening increases with the time width of the incident ultrashort pulses, the overall broadening ratio is low (< 1.02). Because the length of the tapered optical ber probe used is 7 mm and the dispersion and the nonlinearity only vary signi cantly when the cladding diameter is less than 20 µm. Therefore, the time domain broadening of the ultrashort pulses transmitted in the tapered optical bered probe is mainly caused by dispersion. The blue solid and dashed lines in Fig. 6 represent the spectral shift for 780 nm and 1030 nm, respectively. The solid and dashed lines overlap, and both remain at zero as the time domain width of ultra-short pulse increases. When the ultrashort pulse is transmitted in the tapered optical ber probe, the overall nonlinear coe cient of the tapered optical ber probe is low. The tapered optical ber probe does not bring a signi cant spectral shifting to the ultrashort pulses because the interaction length of the ultrashort pulse and the high nonlinearity part of the tapered optical ber probe is limited.
A miniaturized, exible, and robust ber probe is fabricated without bulky optical or mechanical parts. Bene ting from the all-ber design, the ber probe can be recycled after disinfection and sterilization.

Results
We rst demonstrated the capability of distinguishing different Raman spectra of our tapered optical ber probe by using two microspheres samples. One sample was 15-µm polystyrene microspheres (PS, PSMS-1.07, Cospheric), and another sample was 15-µm polymethyl methacrylate microspheres (PMMA, PMPMS-1.2, Cospheric). Figure 7(a) shows the microscopic image of the PS and the PMMA microspheres. Both microspheres were deposited on the slides. The measured CARS spectra were obtained for dark background (the blue square in Fig. 7(a)), PMMA microspheres (the green square in Fig. 7(a)) and PS microsphere (the red square in Fig. 7(a)), and the results are plotted with the same color code as in Fig. 7(b). The Raman resonance peaks at 2912 cm −1 and 3064 cm −1 for the PS microspheres, and 2953 cm −1 for the PMMA microspheres are visible. These experimental results demonstrate the ability of tapered optical ber probes to perform chemical-speci c detection by detecting CARS spectra.
Next, we moved to simulated cells to demonstrate the capability for single-cell endoscopic detection. Here a hollow glass microbubble xed on the hold ber served as simulated cells. A microtube with a diameter of 12 µm was connected to a syringe pump. The microtube was manipulated by a three-dimensions (3D) micromanipulator so that the different liquids simulating cytoplasm could be readily injected into the simulated cell. The tapered optical ber probe mounted on another 3D micromanipulator was carefully inserted into the simulated cell and delivered the excitation lights. The F-CARS signals were collected by the collection ber and detected by the PMT. The physical micrograph of tapered optical ber probe probing a simulated cell is shown in Fig. 8(a). Figure 8(b) and (c) show that the simulated cell was being lled with liquid and had been lled with liquid, respectively.  Figure 9(b) shows the CARS spectra of injected ethanol in the microbubble. Due to the limited resonance spectrum resolution of the CARS system and the weak energy of the Raman resonance peak of ethanol at 2973cm −1 , the Raman resonance peaks at 2927 cm −1 and 2973 cm −1 are displayed in one peak. Figure 9(c) shows the CARS spectra of injected cyclohexane. The rst peak signals correspond to the Raman resonance of cyclohexane at 2853 cm −1 , and the second peak signals contain the Raman resonance at 2923 cm −1 and 2938 cm −1 .
Overall, Fig. 9 shows that the tapered optical ber probe can detect CARS signals of the liquid samples in the simulated cell, demonstrating its applicability for a CARS endoscopic detecting setup. The tapered optical ber probe can measure Raman resonance spectra with a resolution of ~50 cm −1 .
In addition to measuring samples qualitatively in simulated cells, the tapered optical ber probe can also be used for detecting the local concentration of a speci c molecule in a cell, which is helpful for tracking molecule uptake and metabolism. To demonstrate the quantitative detection capability of the tapered optical ber probe, we injected acetone with different concentrations into the simulated cell. The CARS signals strength of CH 3 mode provides, in theory, a handle for detecting acetone concentration. The intensity of the CARS signals corresponding to different acetone concentrations is shown in Fig. 10(a).
We calibrate the CARS intensity measured at 2921cm −1 as a function of acetone concentration in water, and the results are shown in Fig. 10(b). The dependence between CARS signals and acetone concentration is approximately quadratic. Such phenomena are due to the coherence property of CARS, that is, the quadratic dependence of signals intensity on the number of vibrational modes 32 . The insert in Fig. 10(a) shows the CARS spectra of 10 v/v% acetone and water, demonstrating the ability of the tapered optical ber probe to discriminate between CH 3 and OH bonds in the solution.

Conclusion
In conclusion, we proposed and demonstrated a exible, endoscopic, and minimally invasive CARS measurement with a tapered optical ber probe for single-cell application. An adiabatic taper optical ber probe was used to realize CARS excitation light focusing without mismatch at the focal point. An FMF with generated FWM band in 3400 cm −1~3 800 cm −1 , which is out of CARS signals band (2700 cm −1~3 200 cm −1 ), is selected for adiabatic taper optical ber probe fabrication, CARS excitation lights delivery, and CARS signals collection. We theoretically analyzed the fraction of the optical power in the ber, dispersion and nonlinearity of optical ber probe, which indicated that fabricated tapered optical ber probes with 11.61 µm distal cladding diameter enable to deliver and focus the excitation light effectively and cause no signi cant time domain broadening or spectral shift on the excitation pulses.
The measurements for proof-of-concept were made with methanol, ethanol, cyclohexane, and acetone injected into simulated cells. We believe that the ber probe opens up exciting perspectives for intraoperative label-free detecting for real-time histopathology diagnosis.    The FWM signals of MMF cover from ~2700 cm -1 to 2900 cm -1 and interfere with the detection of C-H bond vibrations. The FWM signals of FMF cover from ~3400 cm -1 to 3800 cm -1 and avoid interference to the detection of C-H bond vibrations. The calculated results of the fraction of optical power in the ber η with 4 μm to125 μm cladding diameters. The blue line denotes that the external medium is water, and the red line denotes that the external medium is ATs. The η is 92.4 % at 780 nm and 85.3 % at 1030 nm when the cladding diameter is 12 μm and the external medium is ATs.  The calculated time domain broadening and the spectral shift of ultrashort pulses with different pulse widths transmitted by a 7 mm tapered optical ber probe. The total time domain broadening ratio is < 1.02 for 780 nm pulses and < 1.005 for 1030nm pulses.     CARS spectra of (a) methanol, (b) ethanol, and (c) cyclohexane detected by CARS ber probe.