The capability of in vivo
fluorescence imaging has expanded rapidly over the past few years1–4
. Despite the progress in spatial resolution1
and imaging speed2, 4
, the achievable imaging depth in live samples remains very limited5–7
, which has hindered the progress of many research fields3, 8, 9
. The bottle neck is that only the ballistic component of a light wave has been utilized for imaging, which experiences exponential attenuation due to random scattering in tissues5, 6
. Here we present fluorescence imaging beyond the ballistic regime by combining single cycle pulsed ultrasound modulation and digital optical phase conjugation. We demonstrate a near isotropic 3D localized sound-light interaction zone. With the exceptionally high optical gain provided by the digital optical phase conjugation system, we can deliver sufficient optical power to a focus inside highly scattering media for not only fluorescence imaging but also a variety of linear and nonlinear spectroscopy measurements.
Controlling wave propagation has been an interesting and important subject in many research fields7, 10–18
. In principle, if one can reverse both the propagation direction and wavefront of the optical wave originating from a point (i.e. a guide star) inside turbid media, one can form an optical focus at the original point regardless the thickness of the turbid media, a process known as optical phase conjugation (OPC)11, 12, 19–21
. For imaging, the challenging task is to freely place a guide star at arbitrary locations inside turbid media. Recently, it has been proposed and experimentally demonstrated to use a sound wave to modulate light to create a guide star for OPC22
. As the scattering of sound waves in tissues is negligible in comparison with light23
, the guide star can be placed at depth far beyond the ballistic regime of light. However, for practical fluorescence imaging, there are two remaining challenges. First, sound and light are both propagating waves in tissues. Even with focused ultrasound, their interaction volume is not 3D confined. Second, given a 3D confined interaction volume, the amount of light that is sound modulated within a highly scattering medium is very small. Thus for practical imaging applications in deep tissues, we need tremendous optical gain (> 103
) for the phase conjugation beam, which cannot be readily provided by a conventional phase conjugation system using photorefractive crystals10, 24, 25
Here, we report fluorescence imaging beyond the ballistic regime with < 40 microns spatial resolution. Different from the previous report22
, we use single cycle focused ultrasound pulses and tightly synchronized near-infrared laser pulses to achieve a near isotropic 3D confined interaction volume. The pulsed light and pulsed sound wave are precisely synchronized such that the light wave illuminates the sample only when the single cycle ultrasound pulse propagates to its spatial focus. In such a way, the sound modulation zone is confined in the transverse direction to < 40 microns by the sound focusing element and in the axial direction to < 40 microns by the temporal profile of the single cycle sound pulse convolved with the temporal profile of the laser pulse. To provide sufficient and durable optical power for fluorescence excitation, we employed digital optical phase conjugation19
(DOPC) to perform phase conjugation.
illustrates schematically the operation of the fluorescence imaging system. A high frequency focused ultrasound transducer launches a single cycle pulse into the sample. A short laser pulse illuminates the sample only when the sound pulse travels through its focus. The wavefront of the frequency shifted light is recorded by the DOPC system using heterodyne interferometry. To measure the fluorescence signal, the DOPC system sends out the phase conjugation beam that precisely propagates to the sound focus. A fluorescence detector measures the power of the emitted fluorescence light. To form a fluorescence image, the entire process is repeated as the acoustic focus is raster scanned inside the sample. The experimental setup () is described in Methods.
Figure 1 a Experimental scheme of fluorescence microscopy by single cycle ultrasound pulse guided DOPC. b Experiment setup. λ/2, half wave plate; PBS, polarizing beam splitter; BB, beam block; BS, non-polarizing beam splitter; BE, beam expander; M, mirror; (more ...)
Although without wavefront control the input laser light becomes randomized by scattering, it can still excite fluorescence, resulting in background signals. To measure the background level, we translated the DOPC phase pattern by ~30 pixels both in y and z directions on the SLM (making the DOPC ineffective, see Supplementary Fig. 1 a–c
). Experimentally we measured the fluorescence signals with and without translating the phase pattern on the SLM and the difference between the two signals was used to represent the fluorescence signal at the sound modulation position. We define contrast as the ratio of this signal difference to the background signal, which is shown in the measured images.
To measure the point spread function (PSF) of the system, we dispersed 6 microns diameter fluorescence beads in a 2 mm thick agar slice and sandwiched the fluorescence agar slice between two 2 mm thick scattering tissue phantoms (μs
=6.42/mm, g=0.9306). The details of the phantom are described in the Supplementary discussion
. show the measured PSF with a sampling step size of 15 microns. The data was resampled with bicubic interpolation, as shown in . Gaussian fittings of the cross sections of the PSF () show that the FWHM of the PSF is 38.6±2.8 microns, 37.9±2.3 microns and 263±90 microns (± 95% confidence bound) along y, z and x directions, respectively. The achieved focus to background ratio (FBR) was ~ 3.7 (Supplementary Fig. 1 d–e
). A similar experiment was also performed with fixed rat brain slices as the scattering media, as shown in Supplementary Fig. 2
a Measured transverse PSF through 2 mm thick tissue phantoms (μs=6.42/mm, g=0.9306). b Measured axial PSF. c and d are the corresponding images resampled with bicubic interpolation. e–g Gaussian fitting of the measured PSF.
To verify that the observed fluorescence signals were indeed originating from the ultrasound modulation, we performed a control test by comparing the measurements with and without powering on the amplifier for the ultrasound transducer. We sandwiched a 1 mm thick fluorescence bead agar layer between two 2 mm thick tissue phantoms (g=0.9013, μs
=10.5/mm). As shown in Supplementary Fig. 3
, the signal was gone with the ultrasound transducer disabled.
To demonstrate the fluorescence imaging capability, we used a glass micropipette to manually create an array of 60 microns diameter holes with 120 microns spacing in a 2 mm thick agar slice and injected 6 microns diameter fluorescence beads inside the holes to create a fluorescence pattern. A direct wide field fluorescence image is shown in . The fluorescence hole array was then surrounded with 2 mm thick tissue phantoms (μs
=6.42/mm, g=0.9306). shows the fluorescence image of the hole array with tissue phantoms around it. Due to random scattering, the image diffused to ~ 2 mm in diameter and the structure information was completely lost. We raster scanned (step size 30 microns) the position of the acoustic focus and performed DOPC based fluorescence excitation and the raw data is shown in . The raw data was resampled with bicubic interpolation, as shown in . For comparison, we show the convolution of the measured PSF () with the direct optical image () in . 2D Gaussian fitting for each fluorescence hole is shown in Supplementary Fig. 4
. We also imaged samples, in which the fluorescence features were completely embedded in the middle of a 4 mm thick scattering medium (g=0.9013, μs
=7.09/mm). (Supplementary Fig. 5
Figure 3 a Direct optical imaging of the fluorescence hole array without tissue phantoms. b Direct optical imaging of the fluorescence hole array surrounded by 2 mm thick tissue phantoms (μs=6.42/mm, g=0.9306). c Image acquired with ultrasound pulse guided (more ...)
In our experiments, we achieved < 40 microns lateral resolution with near isotropic 3D confined modulation zone. The dependence of the modulation zone on experimental parameters is analyzed in the Supplementary discussion
. For applications requiring higher spatial resolution, higher frequency ultrasound transducer can be employed to shrink the modulation zone. In the fluorescence imaging experiments, we used one-photon fluorescence excitation, for which the fluorescence excitation is not 3D confined. The background and the out-of-focus excitations reduce the achievable signal to noise ratio (SNR). However, the background could be dramatically reduced by two-photon excitation at the Ti:sapphire wavelength that was employed in this work. In addition, two-photon excitation can further reduce the size of the PSF by ~
due to the square dependence of the fluorescence excitation to light intensity.
In this work, the observed FBR is 1.5–4, a value that needs to be improved for practical imaging applications. Previous studies14, 26
suggest that the achievable FBR is proportional to Npixel
, where Npixel
is the number of independently controlled phase pixels on the SLM and Nmode
is the number of uncorrelated optical modes at the sound modulation zone. An estimation of the theoretical FBR of our system is presented in Supplementary discussion
. By iteratively focusing light into the sound modulation zone via DOPC, we can potentially achieve a much smaller sound light interaction volume, leading to better spatial resolution and higher FBR due to the reduced Nmode
(see Supplementary Discussion
). Employing an SLM with less pixel-to-pixel coupling, higher filling factor and diffraction efficiency, and lower temporal phase fluctuation can potentially improve the FBR by more than one order of magnitude. Moreover, the sound modulation zone can be shrunken by using higher frequency sound transducers, reducing Nmode
and further improving FBR.
In our experiments, we typically acquire 48–96 interferograms and the recording time for one DOPC operation is 1.2–2.4 seconds. We analyzed the SNR's dependence on the camera's parameters in the Supplementary discussion
. Using cameras with higher full well charge capacity and frame rate, we can potentially increase the measurement speed by at least one order of magnitude.
For many in vivo
imaging applications, a transmission configuration is not suitable. However our technique may be extended to measure sound encoded backscattered light27
In conclusion, we report fluorescence imaging beyond the ballistic regime with a 3D confined sound modulation zone, high optical gain and < 40 microns lateral resolution in random scattering media. With the capability of focusing sufficient optical power inside random scattering media, our technique can be used for not only fluorescence imaging but also a variety of linear and nonlinear spectroscopy measurements, which is expected to find numerous important biomedical applications.