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Abstract
A multimodal multiphoton microscopy (MPM) is developed to acquire both two-photon microscopy (2PM) and three-photon microscopy (3PM) signals. A dual-wavelength Er-doped fiber laser is used as the light source, which provides the fundamental pulse at 1580 nm to excite third harmonic generation (THG) and the frequency-doubled pulse at 790 nm to excite intrinsic two-photon excitation fluorescence (TPEF) and second harmonic generation (SHG). Due to their different contrast mechanisms, the TPEF, SHG, and THG images can acquire complementary information about tissues, including cells, collagen fibers, lipids, and interfaces, all label-free. The compact MPM imaging probe is developed using miniature objective lens and a micro-electro-mechanical scanner. Furthermore, the femtosecond laser pulses are delivered by a single mode fiber and the signals are collected by a multimode fiber, which makes the miniaturized MPM directly fiber-coupled, compact, and portable. Design considerations on using the dual excitation wavelengths are discussed. Multimodal and label-free imaging by TPEF, SHG, and THG are demonstrated on biological samples. The miniaturized multimodal MPM is shown to have great potential for label-free imaging of thick and live tissues.
© 2020 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
1. Introduction
Multiphoton microscopy (MPM) is a laser-scanning microscopy technique based on exciting and detecting nonlinear optical signals from tissues and cells [1–3]. It is considered as one of the best noninvasive means of biological imaging microscopy in tissues and live animals [3], and it has been widely used in cancer detection and brain imaging [4,5]. Typical MPM performs two-photon microscopy (2PM) imaging, including two-photon excitation fluorescence (TPEF) [1] and second harmonic generation (SHG) [2]. TPEF is a primary signal in MPM [3], which originates from intrinsic fluorophores in tissues, such as nicotinamide adenine dinucleotide (NADH), flavin adenine dinucleotide (FAD), and elastin, and exogenous fluorophores by staining the tissue with contrast agents [1]. SHG is an intrinsic signal induced by nonlinear scattering which does not involve the absorption of photons. Therefore, SHG imaging does not suffer from photo-toxicity or photo-bleaching, while both of those could affect TPEF imaging in live specimens [6]. However, SHG occurs only in ordered non-centrosymmetric structures, which limits the SHG signal generation mostly to fibrillar collagen and striated muscle myosin in mammalian soft tissues [2]. Currently, three-photon microscopy (3PM) imaging, such as three-photon excitation fluorescence (3PEF) and third harmonic generation (THG), has also been realized. 3PM uses a longer excitation wavelength than 2PM in the near-infrared (NIR) region, which exhibits deeper penetration depth due to reduced scattering. 3PEF has the same origin as TPEF but lower excitation efficiency [7]. THG is another type of label-free imaging and is generated from optical heterogeneities [8], such as at the interface between an aqueous medium and a lipidic, mineralized, or absorbing structure [9]. Since THG does not require the non-centrosymmetric structure as in SHG or fluorescence molecules as in TPEF, it can be used to image structures that otherwise cannot be imaged by 2PM. THG has lower excitation efficiency than TPEF and SHG due to its higher order nonlinear effect. Together, TPEF, SHG and THG can provide multiple complementary contrasts from tissues label-free, which is essential for in vivo imaging on patients and imaging live cells and tissues [3,10].
Multimodal MPM by combining 2PM and 3PM has been shown to be a powerful tool for in vivo imaging and clinical applications. N. Olivier et al. designed a multimodal MPM including TPEF, SHG, and THG for investigating cellular behavior in developmental biology [9]. Using transgenic embryo, TPEF from cell nuclei, SHG from mitotic spindles, and THG from cell walls were imaged and correlated in the mitosis phases of cells. Multimodal MPM combining TPEF, SHG, and THG, was also demonstrated by B. Weigelin et al. to investigate cancer cell invasion [11]. SHG showed collagen fibers and striated muscle myosin. THG detected cells and tissue interfaces, including adipocytes, nerve fibers, and blood vessels. TPEF was excited from fluorescence staining and green florescence protein (GFP), which labeled tumor cytoplasm, tumor nuclei, and blood vessels. Thus, the multimodal MPM allowed a comprehensive study of cancer cell invasion. A multimodal MPM combining SHG and THG was reported by S. Dietzel et al. to characterize microcirculation in animal models [12]. In their work, SHG from striated muscle myosin and collagen fibers in arterioles, and THG from nerve fibers and red blood cells in arterioles and venules, were detected for intravital determination of hemodynamic parameters. Therefore, multimodal MPM with both 2PM and 3PM is a powerful tool for label-free imaging, which can acquire multiple contrasts simultaneously for obtaining comprehensive information about cells and tissues.
In developing the multimodal MPM systems, the laser system design is a major challenge. Two types of laser systems have been used in multimodal MPM. The first type is a single wavelength femtosecond laser in the 1.0-1.7 µm range for exciting 2PM and 3PM signals simultaneously [11,13,14]. Benchtop multimodal MPM systems with excitation wavelength at 1180 nm were developed to investigate bone [11] and skin [13]. Fluorescence staining was required for TPEF imaging due to the long excitation wavelength. The single excitation wavelength was achieved by a Ti:Sapphire laser pumping an optical parametric oscillator (OPO), which was bulky and costly [11,13]. K. Kieu et al. demonstrated a multimodal MPM using Er-doped fiber (EDF) laser with a single excitation wavelength of 1560 nm, where the fiber laser was fiber-coupled to a multiphoton microscope [14]. However, the major limitation of those single wavelength laser systems in the 1.0-1.7 µm range is that intrinsic TPEF signal from tissues such as NADH and elastin cannot be excited by this long wavelength. Since most intrinsic fluorophores in tissue have a two-photon absorption peak at shorter than 900 nm wavelength [15], a wavelength shorter than 900 nm is required in order to excite the intrinsic TPEF signal in tissues. The second type of laser system used in multimodal MPM is dual-wavelength laser source [16–19]. In dual-wavelength laser system, a longer wavelength (>1 µm) excites 3PM signals such as 3PEF and THG, and a shorter wavelength (<1 µm) excites 2PM signals such as TPEF and SHG. Dual-wavelength laser excitation has been achieved by tuning a Ti:Sapphire laser with OPO to two different wavelengths [16–18]. For example, previously reported dual-wavelength sources used tunable wavelengths at 1180-nm for 3PM and 860-nm for 2PM [17]; 1300-nm for 3PM and 920-nm for 2PM [16]; and 1200-nm for 3PM and 800-nm for 2PM [18]. Alternatively, dual-wavelength laser excitation has also been achieved by sequentially using two lasers with different working wavelengths [19]. D. Small et al. reported a multimodal MPM system based on two lasers – a Ti:Sapphire laser at 810 nm for 2PM and a soliton self-frequency shift (SSFS) of EDF laser at 1700nm for 3PM [19]. However, a major challenge of the current dual-wavelength excitation sources (a laser plus OPO or two lasers) is that the whole laser system is highly complicated and bulky, which limits the clinical application of multimodal MPM. Furthermore, the above reported multimodal MPM systems are all carried out on benchtop microscopes.
A compact multimodal MPM is needed in order to translate multimodal MPM imaging into clinical application. Recently, a multimodal MPM endoscopy with a single excitation wavelength at 1700nm was reported by F. Akhoundi et al. [20]. The laser system was designed by applying soliton self-frequency shift (SSFS) to an EDF laser. In the MPM endoscope, a miniature objective lens was attached to and scanned by a piezo electric tube (PZT). The imaging speed was limited to 15 seconds per frame due to the heavy load of the objective lens, which was very slow and not suitable for in vivo imaging. Moreover, the 1700nm wavelength could not excite intrinsic TPEF from tissues. However, an EDF laser has great potential of obtaining multiple different wavelengths. For example, a three-color EDF laser at 775 nm, 864 nm, and 950 nm was reported for TPEF imaging of cells expressing three types of fluorescent proteins [21].
Therefore, the existing multimodal MPM systems still have many limitations. In order to translate the technique to in vivo imaging and clinical applications, several challenges need to be addressed. First, a simpler laser system that can provide dual-wavelength excitation is needed; Second, all label-free imaging of 2PM and 3PM in tissues, including TPEF, SHG, and THG is needed; Third, a compact imaging head with fiber delivery of the laser light and collection of the signals is needed. In this paper, we will address those challenges. In our previous publications [22,23], we have reported a miniature 2PM system using a frequency-doubled EDF laser. In those papers, only the 790 nm beam was utilized while the residual 1580 nm beam was discarded. The system only performed 2PM but there was no 3PM imaging.
In this paper, we report a miniaturized multimodal MPM for acquiring both 2PM and 3PM images. The miniaturized multimodal MPM uses a dual-wavelength EDF laser system, which provides the fundamental pulse at 1580 nm and frequency-doubled pulse at 790 nm. The 1580-nm pulse is used to excite THG, and the 790-nm pulse is used to excite TPEF and SHG. At 790 nm wavelength, intrinsic TPEF signals can be excited from NADH and elastin. The EDF laser system is fiber coupled to an MPM imaging probe via a single mode fiber (SMF) and the frequency-doubling of the pulse is achieved inside the distal end of the probe, which makes the entire system highly compact and portable. Label-free and multimodal imaging by TPEF, SHG, and THG are demonstrated on ex vivo samples of murine and porcine femur bones. Our results demonstrate the capability and potential of the miniaturized multimodal MPM for future clinical applications.
2. System design
2.1 System configuration
Figure 1 shows the schematic diagram of the miniaturized multimodal MPM. It includes an EDF laser source and an MPM imaging probe. The EDF laser source delivers femtosecond pulse at 1580-nm to the miniaturized MPM through a SMF. Through a pigtailed collimator (LPC-02-1550-8/125, OZ Optics), the laser beam is collimated to a parallel beam with a beam diameter ($1/{e^2}$) of 1.4 mm. The 1580 nm pulse is then focused into a periodically poled MgO:LiNbO3 (PPLN, MSHG1550-0.5-0.3, Covesion Ltd.) in the distal end of the miniaturized MPM via an aspheric lens (A375TM-C, f = 7.50 mm, Thorlabs). Maximum conversation efficiency from 1580 to 790 nm is achieved by optimizing the polarization state of the input beam with a half-wave-plate (HWP) and a quarter-wave-plate (QWP). Selection of proper grating period of the PPLN crystal ensures quasi-phase matching (QPM) [24]. Two wavelengths are obtained after the PPLN, which include the frequency-doubled pulse at 790 nm and the residual fundamental pulse at 1580 nm. After the PPLN, the beam is collimated again via an aspheric lens (A375TM-B, f = 7.50 mm, Thorlabs). For 2PM including intrinsic TPEF and SHG imaging, the 790-nm pulse is utilized as the excitation source, selected by a bandpass filter (RG9, Schott) with the transmission window from 700-1050 nm. For 3PM such as THG imaging, the residual 1580-nm pulse is utilized as the excitation source, selected by a longpass filter (FGL1000, Thorlabs) with a passband for >1000 nm. The glass filters RG9 and FGL1000 are switched when performing 2PM and 3PM sequentially. A 2.4-mm-diameter 2D micro-electro-mechanical system (MEMS) mirror (Mirrocle Tech., Inc.) is used to raster scan the beam in X and Y directions. Two achromatic doublets with focal length of 19 mm (AC127-019-A, Thorlabs) and 50 mm (AC127-050-A, Thorlabs) are used to relay the beam and overfill the back aperture of the objective. The objective is a miniature aspherical lens C392TME-A (f = 2.75 mm, NA = 0.64, Thorlabs). The emitted TPEF/SHG/THG signals are collected in the backward direction and separated from the excitation light by a long-pass dichroic mirror with the edge wavelength of 665 nm (FF665-Di02, Semrock). The collected signals are coupled by an achromatic doublet into a multimode fiber (MMF) patch cable (Ø1500 µm, NA = 0.39, Thorlabs), and sent to a photomultiplier tube (PMT) (H9305-03, Hamamatsu). A blocking filter with a passband from 350-650 nm (FF01-680/SP, Semrock) is placed in front of the PMT to remove the residual back-reflected excitation light. Another bandpass in front of the PMT selects the different signals. TPEF signal is selected by a bandpass filter of 550 ± 44 nm (FF01-550/88, Semrock). SHG signal is selected by a bandpass filter of 390 ± 20 nm (FF01-390/40, Semrock). THG signal is selected by a bandpass filter of 520 ± 17.5 nm (FF01-520/35, Semrock). The field of view (FOV) of the MPM system is 140 µm × 220 µm (328 pixels × 512 pixels). The miniaturized microscope is constructed using a 30 mm cage system (Thorlabs, Inc.) which has a cross sectional dimension of ∼40 mm × 40 mm. The setup has an L shape where the two arms are 35 and 15 cm long respectively. The current size of the miniaturized microscope is mainly limited by the cage system. A customized handheld probe can be designed to package the system into a much smaller dimension in the future.
Fig. 1. Schematic of the miniaturized multimodal 2PM and 3PM with a dual-wavelength EDF laser. SMF: single mode fiber; CL: collimating lens; HWP: half-wave plate; QWP: quarter-wave plate; L: lens; SL: scan lens; TL: tube lens; PPLN: periodically poled MgO:LiNbO3 crystal; F: filter; MEMS: micro-electro-mechanical scanning mirror; DM: Dichroic mirror; OBJ: objective lens; MMF: multimode fiber; PMT: photomultiplier tube.
2.2 Design considerations
The spectral relationship among the two excitation wavelengths and the emitted TPEF, SHG and THG signals are illustrated in Fig. 2. The top-right insert in Fig. 2 shows the measured spectrum (ANDO AQ6317B) of the dual-wavelength laser source. The 790-nm pulse is utilized to excite SHG at 395 nm and a broadband TPEF signal ranging from ∼400-650 nm. The 1580-nm pulse is utilized to excite THG at 527 nm. The THG signal is narrow band which can be selected by a bandpass filter and differentiated from the broadband spectrum of TPEF. In the current setup, excitation of 2PM and 3PM are carried out sequentially by switching between the 790 and 1580 nm excitation beams respectively with two filters. Compared to the previously reported dual-wavelength laser systems used in multimodal MPM, our design significantly simplifies the laser system by utilizing both the frequency doubled (790 nm) and the residual fundamental (1580 nm) pulses from the EDF fiber laser.
Fig. 2. Illustration of the spectral relationship of the excitation and emission wavelengths involved in the multimodal 2PM and 3PM imaging. Top left insert shows the energy diagrams of SHG, TPEF, and THG processes. Top right insert shows the measured spectrum of the dual-wavelength EDF laser with frequency doubling.
The characteristics of the dual-wavelength EDF fiber laser system is listed in Table 1. After the PPLN, the power of the 790-nm pulse is ∼61 mW, and the power of the residual 1580-nm pulse is ∼117 mW. The pulsewidth is 80 fs for both the 790-nm and 1580-nm pulses. The short pulsewidth is achieved by using a piece of SMF to compensate the dispersion in the EDF laser system [23]. A −10 dB bandwidth is characterized due to the multiple-peak structure in the spectrum. The −10 dB bandwidth of the 790-nm and 1580-nm pulses are 35 nm and 115 nm, respectively.
Table 1. Specifications of the dual-wavelength EDF laser source.
One challenge of dual-wavelength excitation is the different beam sizes of the two wavelengths, which can affect the optical power throughput due to the limited clear aperture of the MEMS mirror. Inside the PPLN, the frequency-doubled beam should have a smaller beam waist radius than the fundamental beam by a factor of ${1 \mathord{\left/ {\vphantom {1 {\sqrt 2 }}} \right. } {\sqrt 2 }}$ based on nonlinear optics theory [25], such that
where ${w_{0(\omega )}}$ and ${w_{0(2\omega )}}$ are the beam waists of the fundamental and the frequency-doubled beam, respectively. By Gaussian beam propagation, the beam waist ${w^{\prime}_0}$ after the collimation lens can be written as [26] where λ is the wavelength, f is the focal length of the collimation lens. In Eq. (2), ${w^{\prime}_0}$ depends inversely on ${w_0}$ but linearly on λ. Thus, after the collimation lens, the beam waist for the 790 nm beam still has a smaller radius than that of the 1580 nm beam by a factor of ${1 \mathord{\left/ {\vphantom {1 {\sqrt 2 }}} \right. } {\sqrt 2 }}$ Meanwhile, the divergence angle can be written as [27] As the 790 nm and 1580 nm beams propagate in the lens system, the beam size and divergence angle both depend on the wavelength as shown by Eqs. (2) and (4). Figure 3 shows the Gaussian beam width versus the axial distance to the focal plane of the collimation lens for the 1580-nm and 790-nm beams, respectively. The beam radius of the 1580-nm beam (blue line) is $\sqrt 2 $ times larger than that of the 790-nm beam (red line) at the beam waist position and at the far field position. In the experiment, the beam diameters are calculated to be ∼1.46 mm and ∼1.03 mm for the 1580-nm and 790-nm beams, respectively, at the MEMS mirror position. For the MEMS mirror of 2.4-mm diameter at 45° incident angle, the optical power throughput can be estimated as $1 - {e^{ - 2{{(\frac{{1.2 \times \sin {{45}^ \circ }}}{{w_0^{\prime}(z)}})}^2}}}$. For the 1580-nm and 790-nm beams, the throughput is calculated to be 93.3% and 99.6%, respectively. Proper collimation lens needs to be selected in order to ensure that both the 1580 and 790-nm beams can be captured by the MEMS.Fig. 3. Gaussian beam width versus the axial distance to the focal plane of the collimation lens for the 1580-nm (blue) and the 790-nm (red) beams, respectively.
Another challenge of dual-wavelength excitation is the focal shift due to chromatic aberration between the two excitation wavelengths. Numerical simulation of the focal shift is carried out using Zemax and the result is shown in Fig. 4. Figure 4(a) shows the optical layout used in the simulation, including the scan lens, tube lens, and objective. Figure 4(b) shows the focal shift between the design wavelength (790 nm) and other wavelengths for the optical layout in Fig. 4(a). The focal shift induced by the objective alone is also shown in Fig. 4(b) as the dashed line. The focal shift mainly comes from the objective lens. A focal shift of 54 µm is found between the 790 and 1580-nm wavelengths for the current aspherical lens. In experiment, the focal shift between 790-nm and 1580-nm wavelengths is measured to be ∼50 µm by translating a thin slide. In the current experiment, the sample is manually shifted by 50 µm to match the imaging planes for the 2PM and 3PM, respectively. To compensate focal shift, A. Filippi et al. used separate optical paths for adjusting the beam divergence respectively so that the focal planes of 800-nm wavelength and 1200-nm wavelength were overlapped [18]. The drawback of their approach is the increased complexity and dimension of the setup. In the future, the focal shift will be eliminated by designing custom achromatic objective lens using multiple lens elements.
Fig. 4. Zemax simulation of focal shift. (a) Optical layout of the scan lens, tube lens, and objective lens. (b) Simulated focal shift versus wavelength from 400 nm to 1580 nm.
The spot size of the focused beam by the objective lens is simulated by Zemax. The spot diagrams in the center and on the edges of the FOV at 790 nm and 1580 nm are shown in Figs. 5(a) and 5(c) respectively. In the center of the FOV, the spot is tightly focused into a circular pattern. However, on the edges of the FOV, the spots are distorted and enlarged due to spherical and coma aberrations. The root-mean-squared (RMS) spot radius for one-photon signal at 790 nm is shown in Fig. 5(b) at distances varying from the center to the edge of the FOV. Since Zemax does not simulate nonlinear effects, the spot radius for the 2PM signals are calculated as ${1 \mathord{\left/ {\vphantom {1 {\sqrt 2 }}} \right. } {\sqrt 2 }}$ times of that of the one-photon signal, which is shown as the right Y-axis in Fig. 5(b). The RMS spot radius for one-photon signal at 1580 nm is shown in Fig. 5(d). Similarly, the spot radius for the 3PM signal is calculated as ${1 \mathord{\left/ {\vphantom {1 {\sqrt 3 }}} \right. } {\sqrt 3 }}$ times of that of the one-photon signal, and shown in Fig. 5(d). The corresponding radius of the Airy disc, calculated by 0.61λ/NA, is also shown in Figs. 5(b) and 5(d). An imaging system is considered to be diffraction limited if the RMS spot radius is smaller than the Airy disc radius. For this system, only in the central FOV of ± 20 µm and ± 15 µm it satisfies the diffraction-limited criterion for 790 nm and 1580 nm, respectively. The RMS spot radius becomes larger than the Airy disc radius as the FOV increases due to aberrations. Therefore, with a singlet aspherical lens as the objective, the system resolution will be limited by the aberrations introduced by the optics. The resolution can be improved by using a compound objective lens with aberration correction.
Fig. 5. (a) and (c) are the Zemax simulation of the spot diagrams at the center and edges of the FOV at 790 nm and 1580 nm, respectively. Scale bar is 150 µm. (b) and (d) are the RMS spot radius at different distances from the center of the FOV at 790 nm and 1580 nm, respectively. The left and right Y-axes show the RMS spot radius for one-photon and 2PM (or 3PM) signals, respectively. The corresponding radius of the Airy disc is shown as the dashed line.
Currently the 2PM and 3PM images are acquired sequentially. The two excitation wavelengths are selected sequentially by two bandpass filters respectively. To achieve simultaneous 2PM and 3PM imaging, both the 1580 and 790-nm beams will be shined on tissue simultaneously. The TPEF, SHG, and THG signals can be separated into multiple detection channels based on their wavelengths by using dichroic mirrors and bandpass filters on the detection side. For example, four detection channels can be designed, including 395 ± 10 nm for SHG, 405-517 nm for TPEF, 527 ± 10 nm for THG, and 537-650 nm for TPEF. The THG signal will be detected in the 527 ± 10 nm channel but not in the 405-517 nm or the 537-650 nm channel.
2.3 Performance characterization
To characterize the performance of the system, multimodal MPM images from a silicon nano-chip are shown in Fig. 6. The chip contains different nano-structures in the silicon wafer layer whose layout is shown in Fig. 6(a). Here bulk silicon forms the nano-wires and the nano-receivers labeled as part A, part B includes nano-gratings (spacing around 100 nm), and part C is photonic crystal structure which contains dots array (diameter around 180∼250 nm). Figures 6(b)–6(d) show the TPEF, SHG, and THG images, respectively. The pseudocolor used is red for TPEF, green for SHG, and blue for THG. The merged image of the TPEF, SHG, and THG images is shown in Fig. 6(e), which shows good co-registration among the three contrasts. Silicon has been reported to generate TPEF signal [28]. Thus, the nano-wires and part A, B, and C all show TPEF signal in Fig. 6(b). The SHG and THG contrasts in Figs. 6(c) and 6(d) are likely due to surface effect from the silicon wafer.
Fig. 6. (a) Layout of the silicon nano-chip. (b-e) Multimodal MPM imaging of the silicon nano-chip. (b) TPEF image. (c) SHG image. (d) THG image. (e) Merged image. Pseudocolor: TPEF in red, SHG in green, THG in blue. Scale bar is 50 µm.
Figure 7(a) shows the TPEF, SHG, and THG signal intensity, respectively, versus the excitation power on the silicon nano-chip. The signal intensity is calculated by averaging over the pixels which have an intensity value higher than 10% of the maximum intensity. Two linear fitting lines with slope of 2 and 3, respectively, are also shown. As we can see, the TPEF and SHG signals show a quadratic dependence on the excitation power, while the THG signal shows a cubic dependence on the excitation power. This confirms that the detected TPEF and SHG are two-photon effects, and the detected THG is a three-photon effect. The width and depth of the nano-waveguide on the silicon nano-chip are about 500 nm and 220 nm, respectively. The lateral resolution of the miniaturized microscope can be determined by measuring the lateral point spread function using the nano-waveguide. Figure 7(b) shows the TPEF, SHG, and THG intensity line profile with Gaussian fitting across the nano-waveguide. The $1/{e^2}$ spot radius (half width of the spot at the $1/{e^2}$ of the peak intensity) of the Gaussian fit is found to be 0.74 µm, 0.70 µm, and 1.43 µm, for TPEF, SHG, and THG, respectively. Since the actual width of the nano-waveguide is smaller than the measured values, those values could be approximated as the lateral resolution.
Fig. 7. (a) Signal intensity of TPEF, SHG, and THG respectively versus excitation power. (b) Intensity line profile of TPEF, SHG, and THG respectively measured across a nano-waveguide. The dashed lines are the corresponding Gaussian fitting.
In theory, the diffraction limited resolution can be calculated by 0.61λ/NA, which is the radius of the Airy disc based on Rayleigh criterion. This definition applies to one-photon microscopy (1PM). In principle, the resolution of 2PM and 3PM could reach ${1 \mathord{\left/ {\vphantom {1 {\sqrt 2 }}} \right. } {\sqrt 2 }}$ and ${1 \mathord{\left/ {\vphantom {1 {\sqrt 3 }}} \right. } {\sqrt 3 }}$ times, respectively, of the 1PM resolution, because their signal intensities depend quadratically or cubically on the input light intensity [29]. However, this is usually difficult to achieve in experiment due to aberrations [12,30–32]. As shown by Zemax simulation in Fig. 5, only the central FOV is diffraction-limited, while the RMS spot radius are much larger on the edge of the FOV due to aberrations. The experimentally measured resolutions match reasonably well with the simulation in Fig. 5. Table 2 summaries the system performance for the 2PM and 3PM. The resolution of the 3PM is two times worse than that of the 2PM due to the longer wavelength used in 3PM.
Table 2. Comparison of the specifications of the 2PM and 3PM.
3. Multimodal MPM imaging
The capability of the miniaturized multimodal MPM is evaluated by ex vivo imaging of unprocessed (unstained and unfixed) tissue samples. Murine and porcine femur bones are collected within 1 hour following euthanasia and images are taken within 2 hours following euthanasia. For 2PM imaging, the optical power on sample is ∼45 mW at 790-nm wavelength. For 3PM imaging, the optical power on sample is ∼60 mW at 1580-nm wavelength. The 2PM and 3PM imaging are acquired at the frame rate of 0.4 frames/s with the pixel dwell time of 10 µs/pixel. For the images displayed below, the TPEF and SHG images are not averaged, while the THG images are averaged over 5 frames due to their weaker signals.
Figure 8 shows the multimodal MPM imaging of the mid-diaphysis of murine femur bone. The TPEF, SHG, THG, and merged images of the inner surface of the murine femur bone are shown in Figs. 8(a)–8(d), respectively. The TPEF contrast comes from endogenous fluorophores and collagen cross-link [33], the SHG image shows the lamellae collagen fibril bundles [13], and the THG image shows the interfaces in the lacuno-canalicular framework [13,34]. The TPEF and SHG images show similar structures, which was also reported in Ref. [35]. In the merged image in Fig. 8(d), Volkmann's canals can be clearly identified. MPM images of the outer surface of the murine femur bone are shown in Figs. 8(e)–8(h). The lacuno-canalicular framework is clearly observed. Connective tissue on the femur bone surface is also observed. Thus, the TPEF, SHG, and THG images provide complementary information on the porosity and tissue organizations in the bone.
Fig. 8. Multimodal MPM imaging of the inner surface (top) and outer surface (bottom) of murine femur bone. Pseudocolor: TPEF in red, SHG in green, THG in blue. Scale bar is 50 µm.
Figures 9(a)–9(d) show the TPEF, SHG, THG, and merged images, respectively, of the inner surface in the mid-diaphysis of a porcine femur bone. In the bone tissue, the TPEF contrast originates from endogenous fluorophores and collagen cross-link, and the SHG contrast comes from the lamellae collagen fibril bundles. The structure detected by THG is likely lipids [13]. Thus, the multimodal MPM imaging can detect different tissue components and provide complementary information about tissues.
Fig. 9. Multimodal MPM imaging of the inner surface of porcine femur bone. Pseudocolor: TPEF in red, SHG in green, THG in blue. Scale bar is 50 µm.
4. Conclusions
A miniaturized multimodal MPM which can acquire both 2PM and 3PM images has been demonstrated. A dual-wavelength EDF laser generates the fundamental pulse at 1580 nm and the frequency-doubled pulse at 790 nm. The 1580 nm pulse excites THG and the 790 nm pulse excites intrinsic TPEF and SHG, eliminating the need of fluorescence staining. The EDF laser is directly coupled via a SMF to the miniaturized MPM imaging head and the entire MPM system is highly compact and portable.
Our design has made significant improvements on multimodal MPM. First, multimodal 2PM/3PM imaging using a single laser system is achieved. Compared to other dual-wavelength laser systems used in multimodal MPM, our design greatly simplifies the laser system and reduces the cost of the system. Second, all label-free 2PM and 3PM imaging is achieved, including TPEF, SHG, and THG. All label-free imaging is important for in vivo imaging and clinical applications where staining is not possible. Third, our multimodal 2PM/3PM system is compact and all-fiber connected, which makes it portable and can potentially be used in clinic.
Our miniaturized multimodal MPM is capable of acquiring multiple contrast signals, including TPEF, SHG, and THG, in biological tissues all label-free. Complementary information is obtained from intrinsic tissue fluorophores, elastin fibers, collagen fibers, lipids, and interfaces. By adding 3PM imaging to the MPM system, structures that are not visible under 2PM imaging, such as lipids, nerves, adipocytes, and blood vessels can now be seen. The miniaturized multimodal MPM is shown to be a powerful tool for label-free imaging. Its compact design can further enable the translation of the technique into clinical applications.
Funding
Canadian Institutes of Health Research (CPG-151974); Natural Sciences and Engineering Research Council of Canada (CHRP 508405-17); Natural Sciences and Engineering Research Council of Canada (RGPIN-2017-05913).
Acknowledgments
We would like to thank the Centre for Disease Modeling and Animal Care Services at the University of British Columbia for the animal samples.
Disclosures
The authors have no financial interests or potential conflict of interest to disclose concerning this work.
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