Telecommunication technologies

Fig. 4

Gradual beam axial confinement. a Ultrasonic waves (US) create a pressure standing wave (green overlay) that modulates the local refractive index of the tissue, sculpting an optical waveguide. b Schematic representation of the experiment, the dashed black lines indicate the highest and lowest location of the top facets of the tissue phantoms. c Experimentally imaged 2D cross-sections of the optical beam profile propagating inside an agar phantom at different heights along the z-axis, showing an effective and gradual confinement of light from z = −8.1 mm to z = −0.1 mm. d The gradual confinement of the beam. In each plane, the dots represent the boundaries of the full width at half maximum of the beam, while the dashed lines show the extrapolated axial trajectory. The intensity has been normalized to the intensity at z = −0.1 mm case. e Full width at half maximum (FWHM) of the confined optical beam along the axial direction
To visualize the optical beam profile of the sculpted waveguide within the ultrasonic cavity, we imaged the output facet of a series of agar phantom rods of different thicknesses. The phantoms were made of 2% agar at 5 different thicknesses of 2, 4, 6, 8, and 10 mm. The phantoms were placed inside a cylindrical ultrasonic transducer so that the input facet is located 10 mm into the transducer at z = −10 mm (Fig. 4b). Different tissue phantom rods were successively placed inside the transducer and the output facets were imaged from top in the transmission mode (Fig. 4b). We imaged each sample 100 μm below the top surface (output facet) to reconstruct the evolution of the profile (Fig. 4c). The results show that light is gradually confined to a small spot at the output facet at z = −0.1 mm (Fig. 4d). The FWHM of the beam at the input facet (i.e., at z = −8.1 mm) is 200 µm, which is then gradually reduced by ~13 times at the output facet (i.e., z = −0.1 mm), where it measures only 15 µm (Fig. 4e).

Steering light using reconfigurable ultrasonic waves

In addition to piping light deep into scattering media, this acousto-optic method can be used to steer light to target different locations within the tissue. Since light follows the pattern of high-pressure regions defined by ultrasound, we can steer the trajectory of light by simply moving or changing the pattern of ultrasound. Moreover, the pattern of ultrasound is also a function of time, and as a result, by changing the phase difference between the pulsed input laser and the ultrasound wave, we can change the pattern of confined light, since photons experience different patterns of high-pressure regions in the tissue at different relative phases. To create a reconfigurable pattern of ultrasound, we need an ultrasonic phased array, where we can independently control the phase and amplitude of each element. In principle, this phased array can consist of a number of piezoelectric elements spatially arranged in any desired shape around the tissue; this is already done for state-of-the-art ultrasonic medical beam-forming32^,33^,34.
In this paper, we employ an array of 8 piezoelectric transducer elements arranged around a cylindrical cavity (Fig. 5a, r = 19 mm). Each element can be independently addressed electrically. The interference between the ultrasound waves launched from different elements results in a standing wave in the region encompassed by the transducer elements; this can be reconfigured by changing the amplitude and phase of the electrical drive signal to each actuator. To demonstrate the optical mode reconfiguration, we discuss a few examples here. Consider an actuation pattern that produces a dipole mode consisting of two out of phase-split lobes (Fig. 5b). Although these split lobes are spatially symmetric, they are interleaved in time (i.e., they happen at different times that correspond to a π phase difference). If the laser is not pulsed and the image is averaged over many cycles of the ultrasound wave, an interference pattern is formed (Fig. 5c). Numerical simulation of the pressure waves at the top cross-section of the transducer (Fig. 5b) clearly shows a pressure profile consisting of two split regions near the center at two opposite phases; i.e., when one region is at the highest positive pressure and thus forms an optical waveguide, the other is at the highest negative pressure and cannot confine light. These two regions alternate between high pressure and low pressure states every half cycle. Therefore, by synchronizing the pulsed laser with one of the high-pressure regions, we can ensure that light is piped through that region and not the other. By flipping the relative phase between the two sets of elements, light can be switched over to the other waveguide (Fig. 5d, e). In this way, the formation of degenerate modes allows for digital steering of light from one spatial region to another through mode hopping.

Multi-segment transducer array generates complex light patterns. a A multi-segment cylindrical transducer array. Each segment is a cylindrical section with a height of h = 30 mm. b Simulation of the ultrasonic pressure dipole mode obtained within a cylindrical phased array, where opposite phase voltages are applied to sections 1, 2, 3 and 5, 6, 7. The optical image experimentally captured when: c the laser is continuously on (no modulation), d the laser is pulsed and synchronized with the ultrasonic transducer at 46° phase, and e the laser is synchronized with the ultrasonic transducer at 226° phase. The intensity of each 2D cross-section is normalized to its maximum value and the scale bar is 500 μm
We can also form more complex and reconfigurable patterns using the phased array by continuously changing the amplitude and/or phase. In general, our degrees of freedom are: the frequency of ultrasound, the laser pulse modulation, the relative phase between the laser modulation signal and the ultrasonic wave, and the phase and amplitude of each piezoelectric element. By changing the frequency of ultrasound, we can achieve different complex patterns of pressure waves in the region where the launched ultrasound waves from each element interfere and form a complex interference pattern. Also, by tuning the relative phase of the ultrasound excitation and the pulsed laser modulation, we can force light to interact with only a certain part of the ultrasonic standing waves. By changing the amplitude and phase of the piezoelectric elements we can reconfigure the pattern of pressure waves and as a result, the optical beam pattern can be dynamically reconfigured.
As an example, if six elements (1, 2, 3, 5, 6, and 7) of the previously mentioned eight-segment array (same as Fig. 5a) are driven at a much higher frequency of f = 1.216180 MHz to excite higher order interference modes, a complex pattern of ultrasonic standing waves form that results in the optical beam pattern shown in Fig. 6a for a continuous laser. Different parts of the pressure profile are indeed out of phase. As a result, if we modulate the laser at f  = 1.216180 MHz and synchronize it with the ultrasound wave, we can achieve completely different patterns of light by tuning the relative phases. For example, if the laser modulation and the ultrasound waves are synchronized with no phase difference, we can achieve the pattern shown in Fig. 6b, whereas we would achieve the pattern shown in Fig. 6c, if the phase difference is π. Note how the four spots around the center in Fig. 6b disappear and light couples instead to the center spot in Fig. 6c. The pattern in Fig. 6a is the superposition of these two patterns. By continuously changing the phase difference, we can continuously change the optical pattern between that of Fig. 6b, c. We can see from this example that by changing two of the parameters, i.e., the frequency of ultrasound waves to excite higher order modes and the relative phase between the ultrasound waves and laser light, we can sculpt complex patterns of light in the tissue.
Fig. 6

High frequencies excite higher order optical modes. Six elements (1, 2, 3, 5, 6, and 7) of the eight-segment array (same as Fig. 5a) are driven at a higher frequency of f = 1.216180 MHz to excite higher order modes. Experimental images showing that a continuous light (no modulation) and integration over time produces the superposition pattern (a) composed of the two modes in b and c. Note how the four spots around the center in b disappear and light couples instead to the center spot in c. The laser light modulation phase shift is b 0° and c 180° with respect to the ultrasound waves. The intensity of each 2D cross-section is normalized to its maximum value and the scale bar is 500 μm

Ultrasound can guide light in mouse brain tissue

To demonstrate the application of our proposed acousto-optic method for in-tissue light beam guiding, experiments were performed on mouse coronal brain slice tissues (Fig. 7a) (see Methods for details of brain slice preparation). We used a 240 µm thick tissue, which is a typical thickness of brain tissue used in slice electrophysiology. Blue light, as well as higher wavelengths in the visible spectrum, is particularly relevant for optogenetic stimulation protocols (e.g. λ = 480 nm). As wavelength decreases, scattering of light in tissue increases nonlinearly and absorption becomes dramatically larger The scattering and absorption properties of the mouse cortical tissue depends on the type, age, and region of the cortex. Different values in the range of 6.98–29.54 cm⁻¹ at wavelengths λ in the range of 453–480 nm have been reported in literature for the reduced scattering coefficient of a mouse cortical tissue. The reduced scattering coefficient increases as the wavelength is decreased. We used a blue laser at λ = 405 nm as an extreme case, which undergoes a large amount of scattering and absorption in the tissue (Fig. 7b). At this wavelength, the optical thickness for the cortical region of the 240 µm thick mouse brain tissue is larger than 7ℓs. Laser light impinges on the brain tissue from the bottom and the ultrasonic phased array is placed around the brain slice. Using our technique, light can be piped and collected at a spot using a virtual waveguide through the brain slice tissue thickness (Fig. 7c). A line cross-section of the optical waveguide in the brain tissue is shown in Fig. 7d, where it can be clearly seen that light is confined in the core of the waveguide. The measured extinction ratio is ER1~5. The ultrasonic steering of the optical beam at the boundary of different regions in the brain with inhomogeneous scattering and absorption properties would require careful adjustment of the ultrasonic phased array properties to compensate for the effects of the irregularities within the tissue on the ultrasonically sculpted light paths.