Multifocal multiphoton microscopy based on multianode photomultiplier tubes

Abstract

Multifocal multiphoton microscopy (MMM) enhances imaging speed by parallelization. It is not well understood why the imaging depth of MMM is significantly shorter than conventional single-focus multiphoton microscopy (SMM). In this report, we show that the need for spatially resolved detectors in MMM results in a system that is more sensitive to the scattering of emission photons with reduced imaging depth. For imaging depths down to twice the scattering mean free path length of emission photons (2xl (s) (em)), the emission point spread function (PSF(em)) is found to consist of a narrow, diffraction limited distribution from ballistic emission photons and a broad, relatively low amplitude distribution from scattered photons. Since the scattered photon distribution is approximately 100 times wider than that of the unscattered photons at 2xl (s) (em), image contrast and depth are degraded without compromising resolution. To overcome the imaging depth limitation of MMM, we present a new design that replaces CCD cameras with multi-anode photomultiplier tubes (MAPMTs) allowing more efficient collection of scattered emission photons. We demonstrate that MAPMT-based MMM has imaging depth comparable to SMM with equivalent sensitivity by imaging tissue phantoms, ex vivo human skin specimens based on endogenous fluorophores, and green fluorescent protein (GFP) expressing neurons in mouse brain slices.

Fig. 1

An experimental setup for quantifying photon scattering effects on PSF_em. Excitation beam path is very similar to conventional multiphoton microscopy. For measuring the PSF_em FWHM variation, a CCD camera is used as a detector. For measuring the PSF_em spatial distribution, a pinhole aperture of various sizes is positioned in the image plane and a PMT behind the aperture collects the transmitted emission light.

Fig. 2

Variations in FWHM of ${\text{PSF}}_{\text{tot}}^{\text{MMM}}$ as a function of imaging depth due to photon scattering. FWHMs in (a) transverse direction and (b) axial direction were measured from surface down to the imaging depth of $2\times l_{\text{s}}^{\text{em}}(l_{\text{s}}^{\text{em}}=62.5\mu m)$ at emission wavelength (515 nm). There was no significant change of ${\text{PSF}}_{\text{tot}}^{\text{MMM}}$ FWHM in either direction.

Fig. 3

Effects of emission photon scattering on PSF_em peak intensity and shape. (a) Decay profiles of PSF_em peak intensity and integrated intensity (over a 200 μm diameter circular area) as a function of normalized imaging depth down to $2.5\times l_{\text{s}}^{\text{em}}$. The lines denote best exponent fit of the data points. (b) Variations in the PSF_em lateral profile (PSF_em(r, φ) in cylindrical coordinates) with the imaging depth. The PSF_em profile was measured up to 100 μm in radius from peak location. Data points are connected by lines for eye guidance.

Fig. 4

A simulation of photon scattering effect on image contrast in MMM tissue imaging. An original image comprising 5 × 5 microspheres of 10 μm diameter equally spaced with 40 μm separation distance and the experimentally measured PSF_tot was used in the simulation. (a) MMM image at the sample surface, (b) MMM image at imaging depth of $2.5\times l_{\text{s}}^{\text{em}}$, where increased background noise is observed in the center of the image. (c) The decay of SNR of the MMM image at its center location as a function of the imaging depth.

Fig. 5

A schematic of the multifocal multiphoton microscope based on a MAPMT. Excitation beams are depicted in red/orange colors and emission beams are in green color. In this figure, only two beam-lets are ray-traced. The excitation beam is splitted into 8 × 8 beam-lets via a microlens array. Multiple excitation foci (8 × 8) scan the specimen. The emission beam-lets are collected either by a CCD camera or a MAPMT which has 8 × 8 pixels. L1, L2, and L3 are lenses. The transmission of light through the microlens array and high NA objective causes appreciable pulse dispersion which can be corrected by pre-chirping using a pair of prisms (not shown).

Fig. 6

Quantifying signal decay with the increase of imaging depth. The intensity decay coefficient of a MAPMT-based MMM stands between that of a CCD-based MMM and that of a conventional SMM. A simple deconvolution algorithm (DC) is applied to MAPMT-based MMM. The decay coefficient of DC stands close to the that of SMM.

Fig. 7

The effect of emission photon scattering on images acquired with MAPMT-based MMM. (a) An image of microspheres at $1.76\times l_{\text{s}}^{\text{em}}$ deep from the surface in the tissue phantom sample. Ghost images appeared due to the scattered emission photons collected in neighboring pixels of the MAPMT. (b) Image after the deconvolution process. (c) Variation of PSF_em due to emission photon scattering. The increase of intensity in the tail region was due to emission photon scattering. (d) Crosstalk in the MAPMT pixels. It was calculated by integrating the PSF_em over the effective detection area of individual MAPMT pixels.

Fig. 8

Human skin images acquired with MAPMT-based MMM. Epidermal layer of the skin is imaged. From left to right, imaging depth goes deeper from surface. Stratum corneum layer (a), stratum granular (b), and basal layer (c) are shown. Cells are visible based on autofluorescence. Image size: 360 μm × 360 μm, input power: 7 mW per focus, objective: 20×, imaging speed: 2.5 frames/s with 320 × 320 pixels.

Fig. 9

Images of GFP expressing neurons in the ex vivo mouse brain acquired with CCD-based MMM (a-c) and MAPMT-based MMM (d-f) at different depth locations (surface, 30 um, and 75 um deep). For images with MAPMT-based MMM, a deconvolution algorithm was applied to remove the effect of emission photon scattering (g-i). The objectiveused is a 20× water immersion with NA 0.95. The input laser power is 300 mW at 890 nm wavelength. The frame rate is 0.3 frames per second with a frame size of 320 × 320 pixels.