Optical microscopy is an invaluable tool in the biological sciences1,2,3,4,5,6,7,8
as it enables three-dimensional non-invasive in vivo
imaging of the interior of cells and organisms with molecular specificity. Unfortunately optical methods are restricted to an imaging depth of a few scattering mean free path lengths9,10,11
, a severe limitation in many research fields3,12,13
. Recently hybrid techniques9,14,15,16,17,18,19
that combine the deep penetration capability of sound waves and the molecular contrast of light waves have greatly exceeded the depth limitation of pure optical methods. However, at these extended depths the achievable spatial resolution is restricted by the dimensions of the sound focus. Here we present an approach to fundamentally break the resolution limit of hybrid imaging technologies in deep tissue. Through iterative ultrasound guided optical phase conjugation (OPC), we shrink the sound light interaction volume and obtain a drastically sharper optical focus. This technology paves the way for deep-tissue fluorescence microscopy for biological research and medical applications.
The shallow optical penetration depth has restricted many research fields: it has forced biologists to use transparent model organisms, monolayer cell cultures or histological sections of tissue, just to name a few compromises. Consequently a lot of effort was dedicated to push the depth range in optical imaging11,20,21,22,23,24,25,26,27
and recently substantial progress has been reported using hybrid approaches that combine light and sound4,9,17
. Yet there is still a need for a technique that can take full advantage of the wealth of fluorescent labels and provide microscopic resolution at depths of 1 mm in tissues or deeper. For this goal, we need the ability to focus light tightly beyond the ballistic regime at arbitrary locations.
Recently, light focusing deep inside tissues was achieved using ultrasound guided optical phase conjugation14,15
and fluorescence imaging was demonstrated with NIR18
excitation. An ultrasound focus, which experiences much less scattering than light, is used as a source of frequency shifted light that can be recorded and time-reversed using OPC. Similar to other hybrid techniques, however, the resolving power at large depths is determined by the size of the ultrasound focus, resulting in modest spatial resolutions of 30–50 microns18,19
. Further the first demonstrations18,19
lacked sufficient contrast for practical biological imaging.
Here we demonstrate fluorescence microscopy beyond the ballistic regime with a lateral resolution of ~12 microns using iterative ultrasound guided digital OPC. We overcome the sound resolution limit by a factor of three and at the same time increase the focus to background ratio (FBR) fivefold. The principle behind our technique can be explained as follows: after traveling through highly scattering media, the incident light field at the ultrasound focus is completely randomized and unfocused. However, if the light was already pre-focused into the ultrasound focus using OPC, a much more confined sound-light interaction would occur.
Let us assume that the transverse profile of the sound modulation zone and hence the phase conjugation beam at the sound focus is defined as
and that we employ two digital optical phase conjugation (DOPC) systems28
, DOPC1 and DOPC2. DOPC1 first illuminates the sample and the sound modulated light is recorded by DOPC2, which is schematically shown in . DOPC2 then generates a phase conjugation beam that focuses back to the sound focus (). Different from the first illumination, the DOPC2 beam has a focused light distribution
at the sound focus. Therefore the emerging sound modulated light has a new spatial profile
. If we let the two DOPC systems take turns to illuminate the sample and to record the sound modulated light, we can achieve a focus profile
, where N is the iteration number ( ).
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) (more ...)
If we assume a Gaussian profile for
and a strong optical focus (large FBR) for a single OPC operation, the transverse FWHM of the PSF decreases as
. The FBR can be estimated by the number of independently controlled phase pixels Npixel
of the SLM divided by the number of uncorrelated optical modes Nmode
present in the ultrasound focus25,29
. In a 2D approximation, the sound-light interaction area is reciprocally related to N and thus the FBR is expected to increase linearly with N. If the initial focus quality is low (FBR < 5), the dependence of FWHM and FBR on the number of iterations is more complicated. We use numerical simulations to investigate this regime, as described in the Supplementary discussion
For fluorescence imaging, the ultrasound focus was raster scanned through the sample and at each position iterative DOPC was performed. The power of the fluorescence emission for each DOPC excitation was measured and the fluorescent background level was subtracted. The background signal was obtained by lateral translation of the DOPC phase pattern18,26,29
(30 pixels in z and y), which makes the phase conjugation ineffective.