Breaking the spatial resolution barrier via iterative sound-light interaction in deep tissue microscopy

Abstract

Optical microscopy has so far been restricted to superficial layers, leaving many important biological questions unanswered. Random scattering causes the ballistic focus, which is conventionally used for image formation, to decay exponentially with depth. Optical imaging beyond the ballistic regime has been demonstrated by hybrid techniques that combine light with the deeper penetration capability of sound waves. Deep inside highly scattering media, the sound focus dimensions restrict the imaging resolutions. Here we show that by iteratively focusing light into an ultrasound focus via phase conjugation, we can fundamentally overcome this resolution barrier in deep tissues and at the same time increase the focus to background ratio. We demonstrate fluorescence microscopy beyond the ballistic regime of light with a threefold improved resolution and a fivefold increase in contrast. This development opens up practical high resolution fluorescence imaging in deep tissues.

Figure 1

(a–d) Schematic illustration of the iterative focus improvement.(a) The initial incident light field (purple) propagates to the ultrasound focus (yellow circle). A simulated speckle pattern at the sound focus (location marked with the white arrows) is shown in the right inset. A small portion of the input light is frequency shifted (green). (b) In the first DOPC iteration, the green light field is time-reversed and is re-focused into the ultrasound focus, resulting in a more confined sound light interaction (right inset). A portion of the light is frequency shifted (purple). (c) The purple light field is time reversed and is re-focused into the ultrasound zone, further shrinking the sound-light interaction zone. (d) After nine iterations, the time-reversed purple light field results in a much improved focus. (e) Experimental setup; PO, Pockels cell; I, Isolator; BS, beam splitter; AOM, acousto-optical modulator; ND, neutral density filter, BB, beam block; DL, delay line; BE, beam expander; P, polarizer; BP, bandpass filter; L1, f = 35 mm lens; L2, f = 50 mm lens; D, fluorescence detector. Scalebar: 10 microns.

Figure 2

(a) Lateral PSF measurement through 2 mm thick tissue phantoms (μ_s = 7.63 /mm, g factor = 0.9013) for iterations 1, 3, 5, 7, and 9.To normalize the peak intensity, the PSF data sets were multiplied by 6.5, 2, 1.5, and 1.5 for iteration 1, 3, 5, and 7, respectively. (b) Axial PSF measurements for iterations 1, 5, and 9. The PSF data for iteration 1 and 5 was multiplied by 6.5 and 1.5, respectively. (c–e) Gaussian fitting of the measured PSF. (f) Fitted transverse FWHM and simulation (mean values and standard deviation). (g) Fitted axial FWHM and simulation (mean values and standard deviation). (h) Measured focus to background ratio and simulation (mean values and standard deviation). (i) Measured ultrasound modulated light power and simulation (mean values and standard deviation). Scalebar: 10 microns. Colorbar in arbitrary units.

Figure 3

(a) Direct widefield image of the fluorescent structure without tissue phantoms.(b) Direct widefield image of the fluorescent structure surrounded by 2 mm thick tissue phantoms (μ_s = 7.63 /mm, g factor = 0.9013). (c) Image acquired with the first round of ultrasound pulse guided DOPC. (d) Image acquired with five iterations. (e) Image acquired with nine iterations. (f) 2D convolution with a 2D Gaussian function (FWHM: 12 microns). Scalebar: a, b: 100 microns, c: 20 microns. Colorbar in arbitrary units.