Non-degenerate 2-photon excitation in scattering medium for fluorescence microscopy

Mu-Han Yang; Maxim Abashin; Payam A. Saisan; Peifang Tian; Christopher G. L. Ferri; Anna Devor; Yeshaiahu Fainman

doi:10.1364/OE.24.030173

1. Introduction

Two-photon microscopy has had an enormous influence on animal studies of in vivo brain activity providing a tool for high-resolution imaging in live cortical tissue [1–4]. Yet, the majority of 2-photon imaging studies, as of today, have focused on the top ~500 μm of the cerebral cortex due to limited penetration of light into biological tissues. Cerebral neurons, however, are wired in circuits spanning the entire cortical depth (~1 mm in mice) [5], and sampling of activity throughout this depth would be required for reconstruction of circuit dynamics. Therefore, increasing the in vivo penetration depth of microscopic imaging is at the heart of the BRAIN Initiative’s efforts focused on large-scale recording of neuronal activity [6]. Here, we demonstrate that the degree of freedom introduced by non-degenerate 2-photon excitation (ND-2PE), using two independently controlled pulsed laser sources of different photon energies, may provide a number of advantages over the conventional methods promising deeper penetration with higher efficiency of excitation in a scattering medium.

Conventional 2-photon microscopy relies on absorption of two equal energy photons (i.e., degenerate 2PE (D-2PE), Fig. 1(a)). The light source, usually a pulsed femtosecond Ti:Sapphire laser, is tuned within the ~740-1000 nm near infrared (NIR) window. For example, 740 nm is used for imaging of nicotinamide adenine dinucleotide (NADH) [7,8] and 800-1000 nm for imaging of intravascular dextran-conjugated fluorescein isothiocyanate (FITC) [9]. Using even longer excitation wavelengths pushes the depth limit further due to an increase in the photon mean free path [10–12]. A recent study using 1280 nm excitation documented imaging of mouse cortical vasculature in vivo down to ~1.6 mm [11] approximately reaching the fundamental depth limit in scattering tissue [13]. This wavelength, however, lies outside the D-2PE spectrum for most of the fluorophores currently used for brain imaging, thus limiting the approach to a handful of optical probes with far red emission (e.g., Alexa 680).

Fig. 1 Schematic energy diagram demonstrating degenerate and non-degenerate multi-photon absorption of a molecule: (a) degenerate 2-photon excitation (D-2PE); (b) degenerate 3-photon excitation (3-PE); and (c) non-degenerate 2-photon excitation (ND-2PE).

Download Full Size | PDF

Another strategy to increasing imaging depth is the use of higher-order multiphoton excitation. Three-photon imaging relying on absorption of three equal energy photons (i.e., degenerate 3-photon excitation (D-3PE), Fig. 1(b)) at 1700 nm has been recently demonstrated to penetrate down to ~1.3 mm providing images of hippocampal neurons expressing red fluorescent protein [14]. Higher-order multiphoton excitation, however, suffers from very low absorption efficiency quantified by the 3-photon absorption cross-section [15,16], resulting in low emitted photon counts thus slowing down image acquisition. In addition, low fluorescence intensity is a significant drawback for real-time imaging of time-resolved biological processes, e.g., neuronal activity. Furthermore, higher laser power needed to overcome the low D-3PE cross-section may harm tissue resulting from repeated dwelling on the same point as required for imaging of activity over time.

Finally, deeper subcortical imaging can be achieved using more invasive approaches such as microendoscopy [17], by removing the overlaying tissue [18], or with technologies that do not rely on optical focusing but at a price of decreased spatial resolution [19]. However, compromising the health of tissue or spatial resolution might not be an acceptable alternative for most cellular-level neuroscience applications.

To increase penetration depth, while avoiding low 3-photon excitation cross section, we explore ND-2PE – absorption of two photons of different energy [Fig. 1(c)]. Non-degenerate excitation of optical species has a long history in the physical chemistry and microscopy communities [20–22]. Lakowicz et al. demonstrated that the fluorescence intensity of a sample can be enhanced at least 1000-fold through non-degenerate excitation. Furthermore, they identified potential resolution gains as a result of the difference in spot sizes for each laser source. In addition, Lakowicz et al. proposed a method for rejecting background noise by detuning the repetition rate of each light source, resulting in beating of the fluorescence, which can be detected using a lock-in amplifier. Fu et al. subsequently demonstrated imaging using this repetition rate detuning scheme [23]. Cambaliza and Saloma [24], Ibáñez-López et al. [25], Blanca and Saloma [26], Xiao et al. [27], Caballero et al. [28], Wang et al. [29] and Dang et al. [30] have computed increased axial resolution of ND-2PE over D-2PE using various constructed illumination geometries. Kobat et al. [10] demonstrated an increase in axial resolution using ND-2PE via side-by-side displacement of the parallel excitation laser beams. Cambaliza and Saloma [24], Blanca and Saloma [26], Wang et al. [29] predicted and demonstrated that these constructed light techniques suppress background fluorescence generation. Lim and Saloma have predicted that error caused by spherical aberration is reduced with ND-2PE versus D-2PE [31]. Quentmeier et al. [32,33] demonstrated UV-fluorescence lifetime measurements using ND-2PE in the visible spectrum, and Robinson et al. [34] demonstrated fluorescence lifetime imaging using ND-2PE combining visible and NIR wavelengths. Recently, Mahou et al. demonstrated that ND-2PE can be used for simultaneous excitation of four different fluorescent proteins in mouse cerebral cortex [35].

For the goal of deep penetration, ND-2PE holds a promise of attaining higher level of signals characteristic to D-2PE with deeper penetration typical for longer wavelengths of SWIR light. Specifically, we assume that one of the beams is chosen within the most “transparent” part of short-wavelength infrared (SWIR) optical window of ~1300-1400 nm or ~1600-1900 nm identified in a recent study [14,36]. This places the second beam within the ~700-1000 nm NIR wavelength range to supply the required photon energy for ND-2PE of visible emission fluorophores (green to red). The NIR beam will experience higher losses to scattering limiting the achievable power at the focal spot deep in the tissue. However, under ND-2PE, the emitted signal is proportional to the product of the power in each beam. Therefore, increasing the SWIR power will compensate for the NIR power due to scattering losses.

In this manuscript, we investigate ND-2PE fluorescence and show experimentally that its efficiency is comparable with D-2PE. Additionally, we show that for imaging through a scattering medium such as brain tissue, ND-2PE can compensate for loss of the excitation laser intensity caused by tissue scattering. In the following sections, we describe the experiments in support of the characterization of ND-2PE efficiency and its benefits for operation in a scattering medium. Our experimental validation begins with an investigation of the excitation cross-section of ND-2PE in comparison with that of D-2PE in a transparent medium. The experimental setup is described in Section 2.1, followed by a description of experimental data and their analysis in Sections 2.2-2.4. In Section 2.5, we measure the attenuation length of D-2PE and ND-2PE in intralipid with different concentrations. In Section 2.6, we analyze the advantage of ND-2PE for imaging in a scattering medium (approximating the brain tissue) using the ratio of absorption cross-sections for ND-2PE and D-2PE calculated in Sections 2.3 and 2.4 together with the measured attenuation length for the NIR and SWIR beams. Here, we use a model governed by the Beer-Lambert law to predict the penetration depth of ND-2PE where a minimum fluorescence signal can be detected. In the last section, we demonstrate that ND-2PE indeed increases the maximum excitation depth as compared to D-2PE. The degree of improvement will vary depending on the fluorophore properties and choice of the NIR and SWIR wavelengths.

2. Experimental setup and results

2.1 Experiment setup

For our experimental validation, we focused on a well-characterized fluorophore, fluorescein, because it has well understood multiphoton absorption properties [37] and is the basis for novel fluorescent voltage-sensitive probes of neuronal activity [38]. We created an experimental apparatus depicted in Fig. 2. Ultrashort pulses were derived from a Ti:Sapphire laser (Coherent Mira-900 or Coherent Chameleon Ultra II, tuned to 825 nm; further referred to as the NIR beam) and an optical parametric oscillator (APE Optical Parametric Oscillator or Chameleon Compact OPO, tuned to 1315 nm; further referred to as the SWIR beam). Thus, ND-2PE was equivalent to D-2PE at 1013 nm. The NIR beam was separated into two arms by a polarizing beam splitter and a half-wave plate. One of the arms was launched into the OPO. A delay line (Thorlabs LTS150) with a resolution of 2 μm, which is temporally equivalent to 6.6 fs, was used to temporally overlap the NIR and the SWIR laser pulses. This step size was chosen considering the pulse width of the NIR and SWIR beams which were around 146 and 185 fs respectively (i.e., 45 and 60 μm in free space). Several lenses and mirrors in each optical path were used to spatially overlap the NIR and SWIR laser pulses as indicated in Fig. 2. Then the two laser beams were combined by a dichroic mirror (Thorlabs, DMSP 1000). The two overlapping and collinear beams were focused by a microscope objective lens (Newport, 5712-A-H 40X) into the sample, which also collected the emitted fluorescence signal in epi-illumination. In order to detect fluorescence, two dichroic mirrors (Semrock, FF678-Di01-25x36 and FF735-Di01-25x36) and a bandpass filter (FF01-530/11) were placed in front of a photomultiplier tube (PMT, Hamamatsu R3896) with which the fluorescence was detected.

Fig. 2 Experimental setup for demonstration of ND-2PE. PBS, polarization beam splitter; HWP, half wave plate; DM, dichroic mirror; FS, fluorescent sample; 40XOBJ, microscope objective; BPF, band pass filter; PMT, photomultiplier. M, mirror.

Download Full Size | PDF

A computer code (Matlab) was devised to automate the temporal overlap of the NIR and SWIR pulses by scanning the delay line and logging PMT readings at each position. The delay line position was then set to the peak position of the measured cross correlation curve. To achieve spatial overlap of the two beams, we estimated the position of the focal plane for each beam using the theory described in [39], which we used to provide a coarse overlap. The procedure is as follows: We inserted a diffuser filter mounted on a translation stage (Newport, M-561D-XYZ) in the light path near the focal plane of the microscope objective and observed the size and motion of the speckle pattern on a target placed behind the filter, while translating the filter in the axial direction (along the z axis) and in the transverse (xy) plane. When the filter was translated along the z axis, the speckle size was maximized as it approached the focal plane. When the filter was translated in the xy plane, the moving speckle pattern reversed its translation direction while passing through the focal plane. Afterward, a precise overlap of the two beams was achieved by optimizing the defocus and tilt of the SWIR beam to maximize the fluorescence signal.

To calculate the beam radius at the focal spot (i.e., beam waist), we used a beam profiler (Thorlabs, BP209-IR). For each of the beams we obtained the beam size at several axial positions for estimation of the divergence angle behind the objective lens [40]. These measurements were then used to back calculate the position of the focal spot. The calculated beam radius at the focal spot (i.e., beam waist) was $w_{N I R}$ ₌1.32 μm and $w_{S W I R}$ = 2.13 μm for the NIR and SWIR beams, respectively.

2.2 Fluorescence signal under D-2PE and ND-2PE

First, we investigated the dependence of the fluorescence signal on the temporal beam alignment [Fig. 3(a)] and power of each beam [Figs. 3(b) and 3(c)] in a transparent medium (500 μM fluorescein in saline). The temporal alignment was optimized by scanning the optical delay line, and the fluorescence signal was detected by the PMT. Irrespective of the delay line position, we observed a fluorescence signal generated by the D-2PE due to the NIR excitation [Fig. 3(a), “D-2PE”]. No fluorescence signal was generated by the SWIR beam alone because of its insufficient photon energy for D-2PE of fluorescein [37]. An increase in the fluorescence signal was detected as the two beams overlapped indicating the additive effect of the ND-2PE occurring due to the simultaneous absorption of NIR and SWIR photons [Fig. 3(a), “D-2PE + ND-2PE”]. This additive property indicated that, under our experimental conditions, depletion of the NIR beam due to its absorption by the fluorophore in the D-2PE regime was insignificant.

Fig. 3 Experimental demonstration of ND-2PE. (a) dependence of the fluorescence signal on temporal alignment; (b) fluorescence intensity dependence on NIR power (SWIR intensity is fixed at 4.16 × 10²³ photons/cm²s); (c) fluorescence intensity dependence on SWIR power (NIR intensity is fixed at 1.18 × 10²³ photons/cm²s).

Download Full Size | PDF

2.3 Power dependence of ND-2PE and D-2PE on NIR and SWIR beam

Next, we examined the dependence of the fluorescence signal on the power of each beam [Figs. 3(b) and 3(c)]. In the absence of SWIR excitation the signal scaled quadratically with the NIR power (green curve in Fig. 3(b)), which confirms D-2PE of the sample with 825nm excitation [41]. With both lasers illuminating the sample, the fluorescence signal increased with an increase in power of each beam. In Fig. 3(b) the SWIR intensity was held constant and in Fig. 3(c) the NIR intensity was held constant. By subtracting the D-2PE fluorescence contribution [Figs. 3(b) and 3(c), green curves] from the fluorescence measured with simultaneous degenerate and non-degenerate excitation [Figs. 3(b) and 3(c), blue curves] we obtained the ND-2PE fluorescence contribution [Figs. 3(b) and 3(c), red curves]. As expected, the ND-2PE signal increased linearly in both experiments [32]. The measured power dependences in Fig. 3(c) do not start from zero because of the fluorescence due to D-2PE at the fixed NIR power. A cubic dependence of the fluorescence intensity with changing SWIR intensity due to 3-photon excitation was not observed due to insufficient SWIR power (emphasizing the low efficiency of 3PE compared with ND-2PE) [40]. Thus, we have confirmed that our choice of NIR and SWIR wavelengths result in both D-2PE and ND-2PE of our fluorescein sample.

2.4 Cross section of ND-2PE

To quantify the absorption efficiency, we computed the ratio of absorption cross-sections for D-2PE ( $σ_{D}$ ) and ND-2PE ( $σ_{N D}$ ). Toward this end, we started with the expression of fluorescence while assuming non-uniform photon flux within each beam [42]:

F_{D} = K \iint σ_{D} \cdot I_{N I R_D} (t, r, z) \cdot I_{N I R_D} (t, r, z) d V d t,

F_{N D} = K \iint σ_{N D} \cdot I_{N I R_N D} (t, r, z) \cdot I_{S W I R_N D} (t - t_{0}, r - r_{0}, z) d V d t,

where

F_{D}

and

F_{N D}

are the detected fluorescence signal under D-2PE and ND-2PE, respectively; K is a product of the quantum yield of the fluorophore, geometry of the imaging system and the fluorophore concentration and is assumed to be independent of the excitation regime;

I_{N I R_D} (t, r, z)

,

I_{N I R_N D} (t, r, z)

, and

I_{S W I R_N D} (t - t_{0}, r - r_{0}, z)

are the spatiotemporal intensity distributions assuming Gaussian beams in space and Gaussian pulse envelope.

t_{0}

and

r_{0}

are the temporal and spatial offset between NIR beam and IR beam. The intensity distribution then can be expressed as

\begin{matrix} I_{N I R_D} (t, r, z) = I_{N I R_D} [\frac{1}{(\sqrt{π} f τ_{N I R_D})} \exp (- \frac{t^{2}}{τ_{N I R_D}^{2}})] \\ {{[\frac{w_{N I R_D} (0)}{w_{N I R_D} (z)}]}^{2} \exp [- \frac{r^{2}}{w_{N I R_D}^{2} (z)}]}, \end{matrix}

\begin{matrix} I_{N I R_N D} (t, r, z) = I_{N I R_N D} [\frac{1}{(\sqrt{π} f τ_{N I R_N D})} \exp (- \frac{t^{2}}{τ_{N I R_N D}^{2}})] \\ {{[\frac{w_{N I R_N D} (0)}{w_{N I R_N D} (z)}]}^{2} \exp [- \frac{r^{2}}{w_{N I R_N D}^{2} (z)}]}, \end{matrix}

\begin{matrix} I_{S W I R_N D} (t - t_{0}, r - r_{0}, z) = I_{S W I R_N D} {\frac{1}{(\sqrt{π} f τ_{S W I R_N D})} \exp [- \frac{{(t - t_{0})}^{2}}{τ_{S W I R_N D}^{2}}]} \\ {{[\frac{w_{S W I R_N D} (0)}{w_{S W I R_N D} (z)}]}^{2} \exp [- \frac{{(r - r_{0})}^{2}}{w_{S W I R_N D}^{2} (z)}]}, \end{matrix}

where

I_{N I R_D},

I_{N I R_N D},

and

I_{S W I R_N D}

are the average photon flux,

τ_{N I R_D},

τ_{N I R_N D},

and

τ_{S W I R_N D}

are the pulse width of NIR and IR.

w_{N I R_D} (0),

w_{N I R_N D} (0),

and

w_{S W I R_N D} (0)

are the beam radius of NIR and IR beam at focal plane and

w_{N I R_D} (z),

w_{N I R_N D} (z),

and

w_{S W I R_N D} (z)

are the spatial evolution of the beam radius of NIR and IR beam. The average photon flux

I_{N I R_D},

I_{N I R_N D},

and

I_{S W I R_N D}

for each beam was approximated given the average power,

P_{N I R_D},

P_{N I R_N D},

and

P_{S W I R_N D},

photon energy,

ℏ ω_{N I R_D},

ℏ ω_{N I R_N D},

and

ℏ ω_{S W I R_N D},

and the calculated beam radius

w_{N I R_D} (0),

w_{N I R_N D} (0),

and

w_{S W I R_N D} (0)

as

I_{N I R_D} = \frac{P_{N I R_D}}{ℏ ω_{N I R_D}} \cdot \frac{1}{π w_{N I R_D}^{2} (0)},

I_{N I R_N D} = \frac{P_{N I R_N D}}{ℏ ω_{N I R_N D}} \cdot \frac{1}{π w_{N I R_N D}^{2} (0)},

I_{S W I R_N D} = \frac{P_{S W I R_N D}}{ℏ ω_{S W I R_N D}} \cdot \frac{1}{π w_{S W I R_N D}^{2} (0)} .

The spatial evolution of the beam radius of NIR and IR beam can be expressed as

w_{N I R_D} (z) = w_{N I R_D} (0) \sqrt{1 + {[\frac{z λ_{N I R_D}}{π w_{N I R_D}^{2} (0)}]}^{2}},

w_{N I R_N D} (z) = w_{N I R_N D} (0) \sqrt{1 + {[\frac{z λ_{N I R_N D}}{π w_{N I R_N D}^{2} (0)}]}^{2}},

w_{S W I R_N D} (z) = w_{S W I R_N D} (0) \sqrt{1 + {[\frac{z λ_{S W I R_N D}}{π w_{S W I R_N D}^{2} (0)}]}^{2}},

where,

λ_{N I R_D},

λ_{N I R_N D},

λ_{S W I R_N D}

are the pump wavelength for D-2PE and ND-2PE. Because the spatial and temporal components of each laser pulse are assumed to be separable, Eqs. (1a) and 1(b) are simplified to

F_{D} (z) = K \cdot σ_{D} \cdot I_{N I R_D} \cdot I_{N I R_D} \cdot T_{D} \cdot S_{D},

F_{N D} (z) = K \cdot σ_{N D} \cdot I_{N I R_N D} \cdot I_{S W I R_N D} \cdot T_{N D} \cdot S_{N D},

where

T_{D}

and

T_{N D}

are temporal overlap integrals for D-2PE and ND-2PE and

S_{D}

and

S_{N D}

are the spatial overlap integrals for D-2PE and ND-2PE:

\begin{matrix} T_{D} = \int [\frac{1}{(\sqrt{π} f τ_{N I R_D})} \exp (- \frac{t^{2}}{τ_{N I R_D}^{2}})] [\frac{1}{(\sqrt{π} f τ_{N I R_D})} \exp (- \frac{t^{2}}{τ_{N I R_D}^{2}})] d t \\ = \frac{1}{π {(f τ_{N I R_D})}^{2}} \int \exp (- \frac{2 t^{2}}{τ_{N I R_D}^{2}}) d t, \end{matrix}

\begin{matrix} S_{D} = \int {{[\frac{w_{N I R_D} (0)}{w_{N I R_D} (z)}]}^{2} \exp [- \frac{r^{2}}{w_{N I R_D}^{2} (z)}]} \\ {{[\frac{w_{N I R_D} (0)}{w_{N I R_D} (z)}]}^{2} \exp [- \frac{r^{2}}{w_{N I R_D}^{2} (z)}]} d V \\ = {[\frac{w_{N I R_D} (0)}{w_{N I R_D} (z)}]}^{4} \int \exp [- \frac{2 r^{2}}{w_{N I R_D}^{2} (z)}] d V, \end{matrix}

\begin{matrix} T_{N D} = \int [\frac{1}{(\sqrt{π} f τ_{N I R_N D})} \exp (- \frac{t^{2}}{τ_{N I R_N D}^{2}})] [\frac{1}{(\sqrt{π} f τ_{S W I R_N D})} \exp (- \frac{t^{2}}{τ_{S W I R_N D}^{2}})] d t \\ = \frac{1}{(\sqrt{π} f τ_{N I R_N D})} \frac{1}{(\sqrt{π} f τ_{S W I R_N D})} \int \exp [- t^{2} (\frac{1}{τ_{N I R_N D}^{2}} + \frac{1}{τ_{S W I R_N D}^{2}})] d t, \end{matrix}

\begin{matrix} S_{N D} = \int {{[\frac{w_{N I R_N D} (0)}{w_{N I R_N D} (z)}]}^{2} \exp [- \frac{r^{2}}{w_{N I R_N D}^{2} (z)}]} \\ {{[\frac{w_{S W I R_N D} (0)}{w_{S W I R_N D} (z)}]}^{2} \exp [- \frac{r^{2}}{w_{S W I R_N D}^{2} (z)}]} d V \\ = {[\frac{w_{N I R_N D} (0)}{w_{N I R_N D} (z)}]}^{2} {[\frac{w_{S W I R_N D} (0)}{w_{S W I R_N D} (z)}]}^{2} \\ \int \exp {- r^{2} [\frac{1}{w_{N I R_N D}^{2} (z)} + \frac{1}{w_{S W I R_N D}^{2} (z)}]} d V . \end{matrix}

We can then use the ratio of Eqs. (5a) and 5(b) to compute the ratio of the Non-degenerate-to-degenerate absorption cross section at any particular combination of excitation wavelengths:

\frac{F_{N D}}{F_{D}} = \frac{σ_{N D} I_{N I R_N D} I_{S W I R_N D} ξ}{σ_{D} I_{N I R_D} I_{N I R_D}},

where

ξ = \frac{S_{N D} \cdot T_{N D}}{S_{D} \cdot T_{D}}

is an overall correction factor that quantifies the spatio-temporal overlap in the NIR and SWIR beams due to differences in the pulse width and beam waist. In our experiment

I_{N I R_D}

=

I_{N I R_N D}

because the single NIR beam simultaneously contributes to both D-2PE and ND-2PE. Therefore, the expression becomes

\frac{F_{N D}}{F_{D}} = \frac{σ_{N D} I_{S W I R} ξ}{σ_{D} I_{N I R}} .

In this experiment, we measured the beam overlap by z-scanning a diffuser and the result indicated a 5-μm offset in axial axis. Therefore, in order to calculate the overall correction factor

ξ

, the spatial overlap integrals must be computed numerically. Using the measured pulse width, beam radius in Table 1, and the laser power and the florescence intensity shown in the Figs. 3(b) and 3(c), we computed

ξ

= 1.1 and

σ_{N D} / σ_{D} = 0.51 \pm 0.1

for this measurement. This value for the

σ_{N D} / σ_{D}

ratio is specific to the pair of excitation wavelengths, and could be improved by tuning the ND-2PE excitation to better match the resonant absorption. Nevertheless, these results confirm that the ND-2PE cross section was within the same order of D-2PE.

Table 1. Parameters for calculating the excitation cross section ratio of D-2PE and ND-2PE

View Table

2.5 Attenuation length of D-2PE and ND-2PE in scattering medium

Next, we conducted a second separate experiment comparing the fluorescence signal generated using ND-2PE and D-2PE in scattering media. To emulate the scattering properties of biological tissue we used intralipid (Hospira Liposyn II 10%) diluted in distilled water with varying concentrations. The intralipid concentrations in our experiment were 0.5%, 0.75%, 1%, 1.5%, and 2% by volume. Specifically, 1% intralipid in water (by volume) has been used as a standard phantom for studies of light penetration in brain tissue [43,44]. The sample, consisting of 500 µM fluorescein in a glass capillary, was submerged in a cuvette with intralipid solution. With the sample held in place, the intralipid cuvette was scanned along the axial direction such that the excitation lasers had to propagate through ever increasing path lengths within the scattering medium. The fluorescence intensity was measured as a function of depth within the intralipid. The experimental setup for these measurements was similar to that in Fig. 2, but the fluorescence signal was collected from the side of the cuvette at 90 degrees relative to the excitation using a second objective [Fig. 4]. Since the sample remained stationary, the distance between the fluorescein capillary and both microscope objectives was held constant for all data points, and the path length that the emitted light had to propagate through the scattering medium on its way to the detector also remained constant. With this collection scheme, the fluorescence attenuation remains constant for each data point. Thus, any change in fluorescence intensity resulted purely from attenuation of the excitation SWIR and NIR beams on the way to the sample. To verify that both beams focused to the same focal plane, the same alignment optimization methods described in the section 2.1 were applied at every depth. The excitation intensity of the NIR and SWIR beams were chosen such that the fluorescence signal of D-2PE and ND-2PE was the same entering the intralipid (i.e., zero depth). At every depth, fifty data points were collected, and from these the average fluorescence signal and its standard deviation around the mean were tabulated.

Fig. 4 Experimental setup for demonstration of penetration depth of D-2PE and ND-2PE. HWP, half wave plate; DM, dichroic mirror; FS, fluorescent sample; OBJ 1, 40X microscope objective; OBJ 2, 20X microscope objective; BPF, band pass filter; PMT, photomultiplier. The arrow indicated the direction of movement of the cuvette.

Download Full Size | PDF

For every intralipid concentration, the data demonstrated a decay in the fluorescence intensity as a function of increasing depth, i.e., the distance that the excitation beams had to propagate through the scattering medium [Figs. 5(a)-5(e)]. To quantify the attenuation, we modelled this process as a Beer-Lambert Law type decay of the excitation intensity:

I_{N I R} (z) = I_{N I R} (0) \exp (\frac{- z}{α_{N I R}}),

I_{S W I R} (z) = I_{S W I R} (0) \exp (\frac{- z}{α_{S W I R}}),

Here z is depth,

α_{N I R}

and

α_{S W I R}

is the attenuation length of the NIR and SWIR beams, respectively, and I(0) is the laser intensity entering the intralipid. Substituting these equations into the appropriate expressions for the fluorescence intensity we find

I_{D_2 P E} (z) = A I_{N I R}^{2} (0) \exp (\frac{- 2 z}{α_{N I R}}),

I_{N D_2 P E} (z) = A I_{N I R} (0) I_{S W I R} (0) \exp [- z (\frac{1}{α_{N I R}} + \frac{1}{α_{S W I R}})],

where A is a constant which contains absorption cross section, collection efficiency, fluorophore concentration and fluorescence quantum efficiency.

Fig. 5 Relative fluorescence intensity as a function of sample depth within varying concentrations of intralipid: (a) 0.5%, (b) 0.75%, (c) 1%, (d) 1.5% and (e) 2%. (f) Attenuation length as a function of intralipid concentrations.

Download Full Size | PDF

To extract α_D-2PE and α_ND-2PE, we first normalized the fluorescence curves to expresses fluorescence relative to its values at zero depth [Figs. 5(a)-5(e)]. We then fit each curve with Eqs. (10a) and 10(b) for D-2PE and ND-2PE, respectively. The resulting the attenuation lengths for D-2PE and ND-2PE, as well as that of each laser, are plotted as a function of intralipid concentration in Fig. 5(f). For 1% intralipid, we find 50% improvement in the attenuation length of ND-2PE (235 ± 0.8 µm) relative to that of D-2PE (156 ± 0.5 µm).

2.6 Theoretical analysis of the penetration depth with ND-2PE in scattering medium

The results of Section 2.4 show that, with our choice of the NIR and SWIR wavelengths, the excitation cross-section for ND-2PE in the transparent medium is about half of that for D-2PE. However, we anticipate that the efficiency of ND-2PE will be higher compared to that of D-2PE in the scattering medium, due to the longer attenuation length of ND-2PE [Fig. 5(f)]. To investigate this hypothesis, we began with a theoretical calculation. We modeled the beam attenuation with the Beer-Lambert law using the attenuation length of D-2PE and ND-2PE from Fig. 5(f). In order to reveal the effect of ND-2PE, for simplicity, the photon flux of the NIR and SWIR beams were set to be equal for the D-2PE and ND-2PE before entering the scattering medium. The fluorescence signal in the simulation was normalized to the fluorescence signal of D-2PE at the entrance to the scattering medium (i.e., zero depth). The calculated result is shown in Fig. 6, where the blue and green curves represent the fluorescence intensity resulting from D-2PE and ND-2PE, respectively. Within about one attenuation length of the NIR beam, the lower fluorescence signal of ND-2PE is a result of its lower excitation cross-section (Section 2.4). Past this point, the advantage of ND-2PE becomes clear. Since the fluorescence intensity depends quadratically on the intensity of the NIR beam, which experiences higher scattering, the ND-2PE signal begins to exceed that of D-2PE. This is because the SWIR beam is attenuated less during propagation in the scattering medium. At a depth of 700 μm, where the lowest D-2PE signal could be detected experimentally in [14], the D-2PE signal drops by ~19.34 dB (the red dotted line in Fig. 6). For the ND-2PE case, however, the same loss is calculated to occur at 910 μm indicating ~210-μm enhancement of the penetration depth.

Fig. 6 Simulation of the fluorescence intensity under D-2PE (blue) and ND-2PE (green) as a function of depth in the scattering medium.

Download Full Size | PDF

2.7 Experimental validation of penetration depth increase via ND-2PE

To experimentally validate the theoretical analysis in Section 2.6, we returned our setup to the original epi-illumination configuration depicted in Fig. 2. Measurements were performed with a fluorescein sample in 1% intralipid. The sample was submerged at varying depths in the intralipid. At every depth, the fluorescence signal at the PMT was maintained at a signal-to-noise ratio (SNR) of 3dB. We started this experiment at depth zero with the NIR beam alone (D-2PE). The power of the NIR beam was increased with an increasing depth to maintain the 3dB SNR of the signal at the detector. This process was continued until the NIR power was exhausted, which set the excitation depth limit for D-2PE. For greater depths, we kept the NIR power at the maximum and introduced the SWIR beam to produce ND-2PE in addition to D-2PE. The power of the SWIR beam was increased with increasing depth to maintain the 3dB SNR of the signal at the detector. This continued until the available SWIR power was exhausted.

The results of this experiment are presented in Fig. 7. With our particular optical setup, we could maintain the 3dB SNR down to ~550 μm within the intralipid with D-2PE alone. With the combination of D-2PE and ND-2PE we surpassed this barrier - increasing the maximum excitation depth to ~750 μm. Thus, under our specific conditions and target SNR, we achieved a ~200um increase in the excitation depth. We expect that with a more optimal choice of the excitation wavelengths and objectives, an even greater increase in penetration depth can be achieved.

Fig. 7 Experimental demonstration that ND-2PE can excite fluorescence at greater depth.

Download Full Size | PDF

3. Discussion and conclusions

In this study, we demonstrated the advantage of ND-2PE providing greater penetration depth in the scattering medium. Increasing the photon flux of the SWIR beam helps to compensate for the scattered loss of the NIR beam. Ultimately, this effect is limited by wavelength-specific tolerance of the tissue for laser power. Further studies will be required to determine the optimal combination of NIR and SWIR wavelengths and powers for achieving the maximum fluorescence signal while avoiding tissue damage. Additional improvements in the excitation efficiency can be achieved by implementation of Adaptive Optics (AO) to correct the phase distortions experienced by the NIR beam [45–47]. Specifically, the SWIR beam, which can be focused deep inside the tissue, can be used as a reference point (“guiding star”) allowing us to adjust the phase of the NIR beam to achieve the maximum spatial overlap of their focal volumes. An NIR beam has been recently introduced to correct distortions of the focal spot in the conventional (degenerate) 2-photon microscopy [48]. The same correction procedure will also compensate for chromatic aberrations of the objective allowing implementation of standard objectives used in degenerate 2-photon microscopy for non-degenerate excitation.

The current study was limited to a fixed combination of the NIR and SWIR photon energies producing the effect equivalent to excitation at 1013 nm. In the future, efficiency of excitation (i.e., excitation cross-section) can be increased through a systematic search for the most efficient combination of photon energies for a particular fluorophore [20]. The exact combination as well as the degree of improvement will vary among fluorophores depending on their chemical structure.

To summarize, ND-2PE may provide a viable alternative to 3-photon microscopy by circumventing low 3PE cross-section while taking advantage of low scattering SWIR illumination. With the 3PE regime as a reference, we can expect orders of magnitude improvement in the excitation cross-section. Compared to the degenerate 2-photon microscopy, this technology will provide the advantage of greater penetration depth similar to that under 3PE. In the future, we envision a parallel development of non-degenerate 2-photon microscopy instrumentation and molecular engineering of organic molecules and fluorescent proteins tailored for ND-2PE.

Funding

NIH BRAIN Initiative (U01 NS090491, R01NS057198); UCSD Center for Brain Activity Mapping (CBAM); NSF (CBET 1445158, ECCS 1405234, ECCS 1507146); ONR (ECE4509); Cymer Corporation.

Acknowledgments

The authors would like to thank Martin Thunemann and Kivilcim Kiliç for their help with preparation of samples.

References and links

1. K. Svoboda and R. Yasuda, “Principles of two-photon excitation microscopy and its applications to neuroscience,” Neuron 50(6), 823–839 (2006). [CrossRef] [PubMed]

2. J. N. Kerr and W. Denk, “Imaging in vivo: watching the brain in action,” Nat. Rev. Neurosci. 9(3), 195–205 (2008). [CrossRef] [PubMed]

3. W. Denk and K. Svoboda, “Photon upmanship: why multiphoton imaging is more than a gimmick,” Neuron 18(3), 351–357 (1997). [CrossRef] [PubMed]

4. F. Helmchen and W. Denk, “Deep tissue two-photon microscopy,” Nat. Methods 2(12), 932–940 (2005). [CrossRef] [PubMed]

5. T. R. Insel, S. C. Landis, F. S. Collins, and The NIH BRAIN Initiative, “Research priorities,” Science 340(6133), 687–688 (2013). [CrossRef] [PubMed]

6. P. Theer, M. T. Hasan, and W. Denk, “Two-photon imaging to a depth of 1000 μm in living brains by use of a Ti:Al₂O₃ regenerative amplifier,” Opt. Lett. 28(12), 1022–1024 (2003). [CrossRef] [PubMed]

7. K. A. Kasischke, E. M. Lambert, B. Panepento, A. Sun, H. A. Gelbard, R. W. Burgess, T. H. Foster, and M. Nedergaard, “Two-photon NADH imaging exposes boundaries of oxygen diffusion in cortical vascular supply regions,” J. Cereb. Blood Flow Metab. 31(1), 68–81 (2011). [CrossRef] [PubMed]

8. T. Takano, G.-F. Tian, W. Peng, N. Lou, D. Lovatt, A. J. Hansen, K. A. Kasischke, and M. Nedergaard, “Cortical spreading depression causes and coincides with tissue hypoxia,” Nat. Neurosci. 10(6), 754–762 (2007). [CrossRef] [PubMed]

9. K. Nizar, H. Uhlirova, P. Tian, P. A. Saisan, Q. Cheng, L. Reznichenko, K. L. Weldy, T. C. Steed, V. B. Sridhar, C. L. MacDonald, J. Cui, S. L. Gratiy, S. Sakadzić, D. A. Boas, T. I. Beka, G. T. Einevoll, J. Chen, E. Masliah, A. M. Dale, G. A. Silva, and A. Devor, “In vivo stimulus-induced vasodilation occurs without IP3 receptor activation and may precede astrocytic calcium increase,” J. Neurosci. 33(19), 8411–8422 (2013). [CrossRef] [PubMed]

10. D. Kobat, M. E. Durst, N. Nishimura, A. W. Wong, C. B. Schaffer, and C. Xu, “Deep tissue multiphoton microscopy using longer wavelength excitation,” Opt. Express 17(16), 13354–13364 (2009). [CrossRef] [PubMed]

11. D. Kobat, N. G. Horton, and C. Xu, “In vivo two-photon microscopy to 1.6-mm depth in mouse cortex,” J. Biomed. Opt. 16(10), 106014 (2011). [CrossRef] [PubMed]

12. R. Kawakami, K. Sawada, A. Sato, T. Hibi, Y. Kozawa, S. Sato, H. Yokoyama, and T. Nemoto, “Visualizing hippocampal neurons with in vivo two-photon microscopy using a 1030 nm picosecond pulse laser,” Sci. Rep. 3, 1014 (2013). [CrossRef] [PubMed]

13. P. Theer and W. Denk, “On the fundamental imaging-depth limit in two-photon microscopy,” J. Opt. Soc. Am. A 23(12), 3139–3149 (2006). [CrossRef] [PubMed]

14. N. G. Horton, K. Wang, D. Kobat, C. G. Clark, F. W. Wise, C. B. Schaffer, and C. Xu, “In vivo three-photon microscopy of subcortical structures within an intact mouse brain,” Nat. Photonics 7(3), 205–209 (2013). [CrossRef] [PubMed]

15. C. Xu, R. M. Williams, W. Zipfel, and W. W. Webb, “Multiphoton excitation cross-sections of molecular fluorophores,” Bioimaging 4(3), 198–207 (1996). [CrossRef]

16. C. Xu, W. Zipfel, J. B. Shear, R. M. Williams, and W. W. Webb, “Multiphoton fluorescence excitation: new spectral windows for biological nonlinear microscopy,” Proc. Natl. Acad. Sci. U.S.A. 93(20), 10763–10768 (1996). [CrossRef] [PubMed]

17. R. P. Barretto, T. H. Ko, J. C. Jung, T. J. Wang, G. Capps, A. C. Waters, Y. Ziv, A. Attardo, L. Recht, and M. J. Schnitzer, “Time-lapse imaging of disease progression in deep brain areas using fluorescence microendoscopy,” Nat. Med. 17(2), 223–228 (2011). [CrossRef] [PubMed]

18. A. Mizrahi, J. C. Crowley, E. Shtoyerman, and L. C. Katz, “High-resolution in vivo imaging of hippocampal dendrites and spines,” J. Neurosci. 24(13), 3147–3151 (2004). [CrossRef] [PubMed]

19. L. V. Wang and S. Hu, “Photoacoustic tomography: in vivo imaging from organelles to organs,” Science 335(6075), 1458–1462 (2012). [CrossRef] [PubMed]

20. J. M. Hales, D. J. Hagan, E. W. Van Stryland, K. J. Schafer, A. R. Morales, K. D. Belfield, P. Pacher, O. Kwon, E. Zojer, and J.-L. Brédas, “Resonant enhancement of two-photon absorption in substituted fluorene molecules,” J. Chem. Phys. 121(7), 3152–3160 (2004). [CrossRef] [PubMed]

21. J. R. Lakowicz, I. Gryczynski, H. Malak, and Z. Gryczynski, “Two-color two-photon excitation of fluorescence,” Photochem. Photobiol. 64(4), 632–635 (1996). [CrossRef] [PubMed]

22. A. Rapaport, F. Szipöcs, and M. Bass, “Dependence of two-photon-absorption-excited fluorescence on the angle between the linear polarizations of two intersecting beams,” Appl. Phys. Lett. 82(26), 4642–4644 (2003). [CrossRef]

23. D. Fu, T. Ye, T. E. Matthews, G. Yurtsever, and W. S. Warren, “Two-color, two-photon, and excited-state absorption microscopy,” J. Biomed. Opt. 12(5), 054004 (2007). [CrossRef] [PubMed]

24. M. O. Cambaliza and C. Saloma, “Advantages of two-color excitation fluorescence microscopy with two confocal excitation beams,” Opt. Commun. 184(1), 25–35 (2000). [CrossRef]

25. C. Ibáñez-López, I. Escobar, G. Saavedra, and M. Martínez-Corral, “Optical-sectioning improvement in two-color excitation scanning microscopy,” Microsc. Res. Tech. 64(2), 96–102 (2004). [CrossRef] [PubMed]

26. C. M. Blanca and C. Saloma, “Two-color excitation fluorescence microscopy through highly scattering media,” Appl. Opt. 40(16), 2722–2729 (2001). [CrossRef] [PubMed]

27. F. Xiao, G. Wang, and Z. Xu, “Superresolution in two-color excitation fluorescence microscopy,” Opt. Commun. 228(4), 225–230 (2003). [CrossRef]

28. M. T. Caballero, P. Andrés, A. Pons, J. Lancis, and M. Martínez-Corral, “Axial resolution in two-color excitation fluorescence microscopy by phase-only binary apodization,” Opt. Commun. 246(4-6), 313–321 (2005). [CrossRef]

29. C. Wang, L. Qiao, Z. Mao, Y. Cheng, and Z. Xu, “Reduced deep-tissue image degradation in three-dimensional multiphoton microscopy with concentric two-color two-photon fluorescence excitation,” JOSA B 25(6), 976–982 (2008). [CrossRef]

30. S. Deng, L. Liu, G. Wang, R. Li, and Z. Xu, “Three-dimensional superresolution in two-color excitation fluorescence microscopy using theta illumination method,” Opt.-Int. J. Light Electron Opt. 121(8), 726–731 (2010). [CrossRef]

31. M. Lim and C. Saloma, “Primary spherical aberration in two-color (two-photon) excitation fluorescence microscopy with two confocal excitation beams,” Appl. Opt. 42(17), 3398–3406 (2003). [CrossRef] [PubMed]

32. S. Quentmeier, S. Denicke, J.-E. Ehlers, R. A. Niesner, and K.-H. Gericke, “Two-color two-photon excitation using femtosecond laser pulses,” J. Phys. Chem. B 112(18), 5768–5773 (2008). [CrossRef] [PubMed]

33. S. Quentmeier, S. Denicke, and K.-H. Gericke, “Two-color two-photon fluorescence laser scanning microscopy,” J. Fluoresc. 19(6), 1037–1043 (2009). [CrossRef] [PubMed]

34. T. Robinson, P. Valluri, G. Kennedy, A. Sardini, C. Dunsby, M. A. Neil, G. S. Baldwin, P. M. French, and A. J. de Mello, “Analysis of DNA binding and nucleotide flipping kinetics using two-color two-photon fluorescence lifetime imaging microscopy,” Anal. Chem. 86(21), 10732–10740 (2014). [CrossRef] [PubMed]

35. P. Mahou, M. Zimmerley, K. Loulier, K. S. Matho, G. Labroille, X. Morin, W. Supatto, J. Livet, D. Débarre, and E. Beaurepaire, “Multicolor two-photon tissue imaging by wavelength mixing,” Nat. Methods 9(8), 815–818 (2012). [CrossRef] [PubMed]

36. L. Shi, L. A. Sordillo, A. Rodríguez-Contreras, and R. Alfano, “Transmission in near-infrared optical windows for deep brain imaging,” J. Biophotonics 9(1-2), 38–43 (2016). [CrossRef] [PubMed]

37. C. Xu and W. W. Webb, “Measurement of two-photon excitation cross sections of molecular fluorophores with data from 690 to 1050 nm,” JOSA B 13(3), 481–491 (1996). [CrossRef]

38. E. W. Miller, J. Y. Lin, E. P. Frady, P. A. Steinbach, W. B. Kristan Jr, and R. Y. Tsien, “Optically monitoring voltage in neurons by photo-induced electron transfer through molecular wires,” Proc. Natl. Acad. Sci. U.S.A. 109(6), 2114–2119 (2012). [CrossRef] [PubMed]

39. Y. Fainman, J. Shamir, and E. Lenz, “Static and dynamic behavior of speckle patterns described by operator algebra,” Appl. Opt. 20(20), 3526–3538 (1981). [CrossRef] [PubMed]

40. L.-C. Cheng, N. G. Horton, K. Wang, S.-J. Chen, and C. Xu, “Measurements of multiphoton action cross sections for multiphoton microscopy,” Biomed. Opt. Express 5(10), 3427–3433 (2014). [CrossRef] [PubMed]

41. L. Shi, A. Rodríguez-Contreras, and R. R. Alfano, “Gaussian beam in two-photon fluorescence imaging of rat brain microvessel,” J. Biomed. Opt. 19(12), 126006 (2014). [CrossRef] [PubMed]

42. D. Kobat, G. Zhu, and C. Xu, “Background reduction with two-color two-beam multiphoton excitation,” in Biomedical Optics (Optical Society of America, 2008), p. BMF6.

43. S. T. Flock, S. L. Jacques, B. C. Wilson, W. M. Star, and M. J. van Gemert, “Optical properties of Intralipid: a phantom medium for light propagation studies,” Lasers Surg. Med. 12(5), 510–519 (1992). [CrossRef] [PubMed]

44. G. Hong, S. Diao, J. Chang, A. L. Antaris, C. Chen, B. Zhang, S. Zhao, D. N. Atochin, P. L. Huang, K. I. Andreasson, C. J. Kuo, and H. Dai, “Through-skull fluorescence imaging of the brain in a new near-infrared window,” Nat. Photonics 8(9), 723–730 (2014). [CrossRef] [PubMed]

45. C. Wang, R. Liu, D. E. Milkie, W. Sun, Z. Tan, A. Kerlin, T.-W. Chen, D. S. Kim, and N. Ji, “Multiplexed aberration measurement for deep tissue imaging in vivo,” Nat. Methods 11(10), 1037–1040 (2014). [CrossRef] [PubMed]

46. N. Ji, D. E. Milkie, and E. Betzig, “Adaptive optics via pupil segmentation for high-resolution imaging in biological tissues,” Nat. Methods 7(2), 141–147 (2010). [CrossRef] [PubMed]

47. N. Ji, T. R. Sato, and E. Betzig, “Characterization and adaptive optical correction of aberrations during in vivo imaging in the mouse cortex,” Proc. Natl. Acad. Sci. U.S.A. 109(1), 22–27 (2012). [CrossRef] [PubMed]

48. K. Wang, D. E. Milkie, A. Saxena, P. Engerer, T. Misgeld, M. E. Bronner, J. Mumm, and E. Betzig, “Rapid adaptive optical recovery of optimal resolution over large volumes,” Nat. Methods 11(6), 625–628 (2014). [CrossRef] [PubMed]

Parameters	Value
λ_{NIR_D}, λ_{NIR_ND}	825 nm
λ_{SWIR_ND}	1315 nm
τ_{NIR_D}, τ_{NIR_ND}	146 fs
τ_{SWIR_ND}	185 fs
w_{NIR_D}, w_{NIR_ND}	1.32 μm
w_{SWIR_ND}	2.13 μm

Non-degenerate 2-photon excitation in scattering medium for fluorescence microscopy

Abstract

1. Introduction

2. Experimental setup and results

2.1 Experiment setup

2.2 Fluorescence signal under D-2PE and ND-2PE

2.3 Power dependence of ND-2PE and D-2PE on NIR and SWIR beam

2.4 Cross section of ND-2PE

2.5 Attenuation length of D-2PE and ND-2PE in scattering medium

2.6 Theoretical analysis of the penetration depth with ND-2PE in scattering medium

2.7 Experimental validation of penetration depth increase via ND-2PE

3. Discussion and conclusions

Funding

Acknowledgments

References and links

Cited By

Figures (7)

Tables (1)

Equations (23)

Optics Express