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We propose two-photon excited fluorescence (TPEF) microscopy employing a novel phase modulation technique of ultra-broadband laser pulses, which allows the relative excitation of an individual fluorophore with respect to other fluorophores. This technique is based on the generation of multi-wavelength pulse train, which independently interacts with each fluorophore. Our technique is applied to dual-color imaging of cells expressing two types of fluorescent proteins. We achieve the selective excitation of one over the other and vice versa. The product of the maximum contrast ratios exceeds 100. We also demonstrate yielded equal excitation rates in the simultaneous excitation. By the selective excitation of a donor fluorescent protein, fluorescence resonance energy transfer imaging is also achieved.
©2009 Optical Society of America
1. Introduction
Two-photon excited fluorescence (TPEF) microscopy has become a powerful tool for investigating biological phenomena because of its inherent advantages, including three-dimensional resolution without a confocal pinhole, high penetration depth with near-infrared light excitation, less out-of-focus photon-induced damage and photobleaching [1,2]. In particular, multi-color microscopy [3,4], fluorescence resonance energy transfer (FRET) microscopy [5,6], and fluorescence cross-correlation spectroscopy (FCCS) [7] are key techniques for visualizing the movement of biomolecules and their interactions with cellar components in a living cell. By combining these techniques, the dynamic interactions of proteins and sub-cellular structures can be further clarified.
These techniques require distinctive excitation modes, in which various fluorophores are excited either selectively or simultaneously. In FRET microscopy, only a donor fluorophore must be excited without exciting an accepter fluorophore, whereas in the case of FCCS, it is necessary to excite both fluorophores equally. In multi-color microscopy, multi-color images are obtained by exciting all of the fluorophores simultaneously or each fluorophore sequentially. In TPEF microscopy using a narrow-band laser pulse, the switching is performed by tuning the laser wavelength. Of course, fast control techniques are required, since it takes several seconds or more to tune the wavelength of the mode-locked femtosecond lasers that are currently available. Moreover, in the simultaneous excitation, there is often a problem in that the TPEF intensity of one type of fluorophore is either much stronger or much weaker than that of the other fluorophores when cells are labeled with various fluorophores with large differences in concentration. Thus, it would be advantageous if the TPEF intensities of the various fluorophores could be controlled individually and arbitrarily in such a situation.
Compared to the use of narrowband pulses, broadband pulses offer the potential for easy and rapid switching between excitation modes, together with the ability to control the respective TPEF intensities from various fluorophores more readily. Assuming that no intermediate resonant level is present, the TPEF intensity I F is proportional to the overlap integral between the two-photon excitation (TPE) spectrum g (2)(ω) of the fluorophore and the second harmonic (SH) power spectrum |E (2)(ω)|2, and is described by
The SH power spectrum is expressed by [8–10]
where |E(ω)| and ϕ(ω) are the spectral amplitude and the spectral phase of an excitation pulse, respectively. Since |E (2)(ω)|2 can be shaped by modulating the spectral phase of an excitation pulse, the TPEF intensity can be controlled by the use of spectral phase modulation. The selective excitation of fluorophores with different TPE spectra is achieved by sinusoidal phase modulation [8–10] and binary phase modulation [11,12]. Joffre’s and Dantus’s groups have demonstrated that sinusoidal phase modulation [13] and binary phase modulation [14] allowed the selective excitation in practical TPEF microscopy. However, it has not reported that the spectral phase modulation provides control of TPEF intensities from various fluorophores individually. This is because the pulses produced by these phase modulation techniques cannot interact independently with each of the fluorophores.
Here, we propose the use of multi-level phase modulation to bring about the above-mentioned potential of broadband pulses. This technique is based on the use of a constructive interference phase to maximize the SH intensity at all frequencies and a destructive interference phase to minimize SH intensity, as obtained by an adaptive control method. The selective excitation can be achieved by the modulated pulses, which are generated only by applying a spectral phase combining the two phase functions. Individual TPEF intensities from various fluorophores can be freely controlled by multi-wavelength pulse train, which is produced by introducing a group delay in a specific spectral region and by using appropriate parameters to determine the hybridization ratio of the two phase functions. Motzkus’s group has recently proposed a different but similar phase modulation technique to generate multi-wavelength pulse train [15]. This technique is suggested for the production of pump and probe pulses in time-resolved coherent anti-Stokes Raman scattering spectroscopy. We demonstrate that the proposed phase modulation technique can be applied to the multifarious control of TPEF intensities from various fluorescent proteins.
2. Design of spectral phase functions
2.1 Spectral phase functions for selective excitation
We propose a spectral phase modulation technique, which enables the selective excitation of a specific fluorophore over other fluorophores that are present. We define constructive ϕ c(ω) and destructive ϕ d(ω) interference phases, which maximize and minimize the SH intensities respectively, over the whole frequency range. ϕ c(ω) generates a Fourier-transform limited (FTL) pulse. The conditions resulting in destructive interference for our technique is similar to that for the binary phase modulation [11]. By designing the fundamental field at jth pixel of a phase modulator to be |E(ωj)| exp{iϕ((ωj)}, we show the conditions resulting in destructive interference. The intensity of the SHG signal at 2ωj is proportional to the square of the superposition of the second harmonic fields produced by the combination of the fundamental field of ωj+ωk and that of ωj-ωk and is expressed by
where the integral in Eq. (2) is now replaced by a discrete sum. The total SHG intensity over whole frequency range is described by
ϕ(ωj) to minimize Eq. (4) is the destructive interference phase. The destructive interference phase depends on the fundamental spectrum. Because it is difficult to obtain an analytical solution, we determine the destructive interference phase by the use of an adaptive control technique employing a simulated annealing (SA) method or genetic algorithm method. The broad SH spectrum obtained by the FTL pulse is divided into two spectral bands (R1 and R2) for two specific fluorophores. R1 gives a high TPE cross-section for one of the fluorophores, whereas R2 gives a high TPE cross-section for the other. To increase the SH intensity within R1 while decreasing the SH intensity within R2, ϕ c(ω) and ϕ d (ω) are applied to R1 and the R2, respectively. The phase function for the selective excitation is expressed by
2.2 Spectral phase functions for simultaneous excitation with various signal ratios
For the simultaneous excitation of two fluorophores that exhibit different TPE spectra, we design spectral phase functions in which ϕ d(ω) is combined with an arbitrary percentage (α or β) of ϕ c(ω) within R1 or R2, respectively. Furthermore, we introduce a group delay phase φ′(ω-ω 0) to the spectral phase function of one region (usually R1) such that the two excitation pulses composed of fundamental spectral components in R1 and R2 interact independently with each of the fluorophores. Individual TPEF intensity is determined by α or β. The spectral phase is described as
3. Experimental
3.1 Experimental setup
The experimental setup [16] is illustrated in Fig. 1. As an ultra-broadband laser, we employed a Ti:sapphire laser (Nanolayers, Venteon) operating at a repetition rate of 80 MHz. The laser spectrum ranged from 660 to 1,100 nm. To compensate for the second-order dispersion of all of the optical components before the focal point of the microscope objectives (OB1 and OB2) (Olympus Corporation, UPLSAPO40×, NA 0.9), whose chromatic aberration and spherical aberration were optimized for the spectral region ranging from the visible region to the near-infrared region, the laser pulses were passed through a sequence of fused silica prisms. A grating-pair-formed pulse shaper with a liquid-crystal spatial light modulator (SLM) (Cambridge Research & Instrumentation, Inc., LC-SLM-128) was employed to compensate for the higher-order dispersion and additional spectral phase modulation. To improve the spectral resolution of the spectral phase modulation, the spectral bandwidths were limited to 340 nm at the SLM, respectively. The beam from the pulse shaper was split into two by an ultra-broadband 50:50 beam-splitter (BS), which gives the same dispersion for both transmission and reflection in the spectral range from 600 to 1,500 nm [17]. Then the two beams were focused onto two different fluorescent protein samples. As the fluorescent proteins, we used pairs consisting of blue- and green-emitting fluorescent proteins (BFP and GFP) or cyan- and yellow-emitting fluorescent proteins (CFP and YFP). The spectra at the sample for a pair of BFP and GFP or a pair of CFP and YFP are shown in Fig. 1(c). Each resultant TPEF signal was detected with a photomultiplier tube (PMT) (Hamamatsu Photonics, R7400U-06). To exclude the excitation wavelengths, short-pass filters (Showa Optronics Co., Ltd., Broadband dichroic filter for Ti:sapphire laser) were placed in front of PMTs. In addition, band-pass filters labeled BP1 (Semrock, Inc., FF02-447/60-25 for BFP or FF01-483/32-25 for CFP) and BP2 (Semrock, Inc., FF01-520/35-25 for GFP or FF01-542/27-25 for YFP) were placed in front of PMT1 and PMT2 respectively to select the detection wavelength. In the TPEF imaging, our experimental setup in Fig. 1(a) was replaced by the setup shown in Fig. 1(b). The sample was scanned by a three-axis piezoelectric transducer (PZT) stage (TRITOR 101 CAP, Piezosystem Jena, Inc.). The input power at the sample was 3.2 mW. The exposure time was set to 200 µs/pixel.
3.2 Multifarious control of TPEF intensities
We first obtained the functions of ϕ c(ω) and ϕ d(ω). One of the output beams was focused onto a 10-µm-thick β-barium borate (BBO) crystal, and two phase functions which maximized and minimized the SH intensities were identified by an adaptive control technique employing an SA method. The measured maximized and minimized SH spectra are shown by the black and red lines in Fig. 2(a), respectively. ϕ c(ω) and ϕ d(ω) are illustrated by the black and red lines in Fig. 2(b) where we removed the phase to compensate for the dispersion of the optical setup.
By combining the obtained ϕ c(ω) and ϕd(ω) functions, we attempted the selective excitation of a BFP over an enhanced GFP and vice versa. Taking into account that the GFP and BFP have peaks at 920 nm [18] and 760 nm [19], in their TPE spectra respectively, the boundary frequency (ω b) between R1 and R2 was varied accordingly. The TPEF intensities of GFP and BFP measured at various ω b values were normalized against those obtained by the FTL pulse, and then plotted as the closed squares in Fig. 3(a). When ω b was 2.30×1015 rad/s, corresponding to a wavelength of 820 nm, the highest contrast ratios of 21:1 and 1:5.6 were achieved for GFP/BFP and BFP/GFP, respectively. The two SH spectra and the spectral phases at the highest contrast ratios are shown in Fig. 2(a) and in Fig. 3(b), respectively. Since the product of the two highest contrast ratios is over 100, this method is applicable to dual-color imaging using BFP and GFP.
Fig. 1. Experimental setup in vitro (a) and in vivo (b). (c) Spectra of broadband pulse at the focal point for a pair of BFP and GFP (black line) and a pair of CFP and YFP (pink line). SLM: spatial light modulator, BS: beam splitter, OB: objective lens, SP: short-pass filter, BP: band-pass filter, PMT: photomultiplier tube.
Fig. 2. (a) SH spectra at constructive (black line) and destructive (red line) interference phases and spectral phases for selective excitations of GFP (green line) or BFP (blue line). (b) Constructive (black) and destructive (red) interference phases.
By mixing the obtained ϕ c(ω) and ϕ d(ω) functions with an appropriate group delay, we achieved simultaneous excitation, together with control, of the respective TPEF intensities from BFP and GFP. The group delay was set to 100 fs in R1, where ω b was 2.30×1015 rad/s. One of the percentage parameters (α, β) in the two spectral regions was set to 0.0 and other was set to 0.0, 0.2, 0.3 and 0.4. The black line and the other color lines in Fig. 3(c) illustrate the applied group delay phase and the spectral phases which were the phases before adding the group phase for controlling the TPEF intensity of BFP while keeping GFP intensity constant, respectively. Figures 3(d) and 3(e) show the SH spectra at the spectral phase used for controlling the TPEF intensities of either one of BFP or GFP while keeping the other intensity constant. The double digits indicate the percentage parameters in the two spectral regions. The SH intensities independently decreased as the percentage parameters increased.
We found that the application of multi-level phase modulation enables individual adjustment of the intensities in two spectral regions in the SH spectrum. The blue and green crosses in Fig. 3(a) indicate the cases where the TPEF intensity of either GFP or BFP is controlled while the other is frozen, respectively. It should be noted that the TPEF intensities could be freely controlled by the use of multi-level phase modulation.
Fig. 3. Control of TPEF intensities from two fluorescent proteins by spectral phase modulation. (a) TPEF intensities from GFP and BFP with selective excitations of GFP (green square) or BFP (blue square) and with simultaneous excitation together with control of TPEF intensity for only GFP (blue cross) or BFP (green cross). (b) Spectral phases at the highest contrast ratio in the selective excitation between a pair of BFP and GFP. (c) Spectral phases in the simultaneous excitation for controlling TPEF intensity of BFP while keeping that of GFP constant. (d) SH spectra at spectral phase for controlling TPEF intensity of BFP while keeping that of GFP constant. (e) SH spectra at spectral phase for controlling TPEF intensity of GFP while keeping that of BFP constant. (f) TPEF intensities from CFP and YFP with selective excitation of CFP (cyan square) or YFP (yellow square) and with simultaneous excitation together with control of TPEF intensity for only YFP (cyan cross) or CFP (yellow cross).
Next, the multi-level phase modulation technique was also applied to the excitation of another FRET pair [20] consisting of an enhanced CFP and a YFP (Venus [21]). Figure 3(f) shows the relationship between the TPEF intensities of YFP and CFP. The orange and cyan squares in Fig. 3(f) indicate the cases where YFP and CFP are enhanced, respectively. The orange crosses in Fig. 3(f) indicate the cases where only CFP or YFP is controlled while the other is frozen, respectively. Each TPEF intensity was normalized against that obtained by the FTL pulse. The highest contrast ratios were 17:1 and 1:5.9 for YFP/CFP and CFP/YFP at ω b of 2.03×1015 rad/s (929 nm) and 2.06×1015 rad/s (913 nm), respectively.
3.3 Dual-color imaging
We applied multi-level phase modulation to the dual-color imaging of a HeLa cell employing BFP (Azurite [22]) and GFP, which were used to label the nucleus and cellular cytoplasm, respectively. By adopting selective excitation with the highest contrast ratios, dual-color TPEF images could be acquired sequentially as shown in Fig. 4(a). Because there is no cross-excitation, this method can also be applied to FRET imaging [20]. Under simultaneous excitation, we obtained dual-color TPEF images (from left-to-right panels) with gradually decreasing intensities for only BFP and for only GFP as shown in Figs. 4(b) and 4(c), respectively. The double digits in parentheses indicate the percentage parameters (α, β) in the two spectral regions. By using variable percentage parameters, the respective TPEF intensities could be freely controlled. Therefore, multifarious control of TPEF intensities was successfully demonstrated. It must be emphasized that only a single light source was used here. Even if there are large differences in concentration among the various fluorophores, this control technique can be used to adjust the TPEF intensities to the same level for multi-color imaging.
Fig. 4. Dual-color images of a HeLa cell labeled with BFP (nucleus, top) and GFP (cytoplasm, bottom). The scale bar is 10 µm. (a) BFP-enhanced excitation (left) and GFP-enhanced excitation (right). (b, c) Simultaneous excitation with control of TPEF intensities. (b) BFP intensity is regulated, while GFP intensity is frozen. (c) GFP intensity is regulated, while BFP intensity is frozen.
3.4 FRET imaging
With CFP-enhanced excitation, FRET imaging can be demonstrated. Yellow-cameleon (YC3.60) was used to label the cellular cytoplasm in a HeLa cell. Cameleon, which has a CFP and a YFP, is an indicator for Ca2+ [23]. Figures 5(a) and 5(b) show TPEF images of CFP and YFP before Ca2+ stimulus excitation by treatment with 5 µM ionomycin, respectively. Figures 5(d) and 5(e) show TPEF images of CFP and YFP after the Ca2+ stimulus excitation, respectively. To eliminate emission crosstalk for CFP and excitation crosstalk for YFP, the net FRET signal (nF) was calculated as follows [24]
where I FRET, I CFP and I YFP are the intensities detected by the YFP channel under CFP-enhanced excitation, by the CFP channel under CFP-enhanced excitation, and by the YFP channel under YFP-enhanced excitation, respectively. a and b, which are the percentage norms for the percentage of CFP and YFP crosstalk in FRET imaging, were 0.064 and 0.12 in our system. FRET images before and after the Ca2+ stimulus excitation are shown in Figs. 5(c) and 5(f), respectively. We found that the FRET signal increased after Ca2+ stimulus excitation due to an increase in the Ca2+ concentration in the cell. We also measured a histamine-induced Ca2+ response in a HeLa cell. Figure 5(g) shows the signal ratio of the YFP channel to the CFP channel before and after Ca2+ stimulus excitation by treatment with 20 µM histamine, monitored every 1 s. We could observe Ca2+ oscillation with a high dynamic range. It should be noted that FRET imaging could be successfully achieved with CFP-enhanced excitation, even if ultra-broadband pulses were used as an excitation light source.
Fig. 5. Ca2+ response in cytoplasm of a HeLa cell loaded with yellow-cameleon (YC3.60). CFP (a, d), YFP (b, e) and FRET (c, f) images with CFP-enhanced excitation before (a, b, c) and after (d, e, f) Ca2+ stimulus excitation by treatment with 5 µM ionomycin. The scale bar is 15 µm. (g) Signal ratio of YFP channel to CFP channel before and after Ca2+ stimulus excitation by treatment with 20 µM histamine.
4. Discussion and conclusions
We demonstrated that TPEF microscopy employing spectral phase modulation could provide rapid switching between selective excitation with a high contrast ratio, and simultaneous excitation with control of the TPEF intensities from some fluorophores with different TPE spectra independently. The selective excitation was applied to FRET imaging. Under a large concentration difference between fluorophores in multi-spectral imaging using a single laser, the TPEF intensity of a specific fluorophore is substantially different from that of other fluorophores. Even if the sensitivities of the detectors are individually adjusted in order to acquire equal signal levels, a strong signal might be detected from one fluorophore by another detector with a high sensitivity for weak signals from other fluorophores due to the overlap between the tail of the fluorescence spectrum of one fluorophore and the peak of the fluorescence spectrum of the other fluorophore. Weak TPEF signals are often buried in the crosstalk Thus, it has previously been necessary to radically adjust the concentrations of fluorophores when preparing samples labeled with various fluorophores. Of course, in the case of the use of multi-lasers, it might not been necessary. However, ultrashort pulse lasers are very expensive. The proposed multi-level phase modulation technique makes the TPEF intensities from the different fluorophores with different TPE spectra equal and allows the reduction of crosstalk. Therefore, it is not necessary to significantly adjust the fluorophore concentrations during the preparation of samples labeled by different fluorophores with different TPE spectra. Recently, Rabitz’s group has demonstrated the selective excitation of fluorophores that exhibit almost identical spectra [25]. Thus, it will be possible for the development of the spectral phase modulation technique to control TPEF intensities from various fluorophores.
In FCCS, the center wavelength of a narrowband pulse has been tuned to the optimal wavelength in order to achieve equal TPEF intensities for two fluorophores. The optimal wavelength corresponds to the cross-over point of two TPE spectra, where the TPE coefficients are often low [7]. Thus, it is necessary to increase the incident intensity in order to obtain a high signal-to-noise ratio, and there is a high probability that increasing the incident intensity will induce photobleaching. This problem can be solved by the use of multi-level phase modulation, in which each fluorophore is individually excited at the wavelength where the TPE coefficient is high. Therefore, photobleaching will be reduced if the photobleaching is proportional to the square of the incident intensity.
Inter-frame spacing is determined by summation of the imaging time and the phase modulation time. In these experiments, it takes 100 ms to modulate the spectral phase. This time, which is around 100 times less than the laser tuning time, is limited by the response time of the liquid crystal SLM. The modulation speed can be improved by selecting a modulator with a short response time. A variety of techniques for spectral phase modulation have been demonstrated to date, including the use of a liquid-crystal SLM [26], a deformable mirror [27], acousto-optic modulators [28] and electro-optical modulators [29]. Among these experiments, electro-optical modulators have been shown to provide modulation speeds of the order of 10 ns [29]. Since this modulation rate is much higher than the frame-rate in real-time imaging [5], the excitation mode can be switched between selective excitation and simultaneous excitation for each frame without dropping the frame-rate. In addition, the TPEF intensities can be independently and arbitrarily controlled in each frame. Even if the concentrations of the various fluorophores are independently changed along with alterations in cellular morphology, we can correct the TPEF intensities by the use of intensity control for each frame. The proposed multi-level phase modulation is an important tool for rapid and easy switching among multi-spectral imaging, FRET imaging and FCCS.
The general low average power of the ultra-broadband pulse laser oscillators limits the imaging depth. The problem about the low power would be solved by high-power ultrabroadband pulse lasers, which is based on the cavity dump technique [30] and/or on the amplification technique [31]. Even though these high-power lasers are low-repetition rate, real-time deep imaging would be achieved by the high-power lasers combined with parallel excitation technique with many foci [32,33].
In conclusion, the multi-level phase modulation demonstrated here allows not only selective excitation with a high contrast ratio, but also simultaneous excitation together with the control of TPEF intensities from different fluorophores with different TPE spectra independently and arbitrarily. Because the spectral phases can be exchanged rapidly, this technology will facilitate TPEF microscopy aimed at clarifying the dynamic interactions of proteins and sub-cellular structures.
Acknowledgments
We thank Takako Kogure, Tetsuya Kitaguchi, and Yoshiko Wada from RIKEN for providing the biological sample. This research was supported by the Special Postdoctoral Researchers Program of RIKEN. This work was partly supported by a Grand-in-Aid for Scientific Research (No. 18656023) from the Ministry of Education, Culture, Sports, Science and Technology, Japan.
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