# Wavefront Shaping and Its Application to Enhance Photoacoustic Imaging
## 1 Introduction  
High-resolution optical imaging has recently seen more and more biomedical applications, such as optical coherence tomography (OCT) for optometry [1], confocal microscopy for cytobiology research [2], two- or three-photon microscopy for deep tissue imaging [3; 4; 5; 6; 7; 8], as well as stimulated emission depletion (STED) microscopy and photo-activated localization microscopy (PALM) for super-resolution imaging of single molecules [9; 10]. In almost all of these applications, photons, whether illuminating or excited ones, need to propagate some distance in tissue. Therefore, the capability of focusing light plays an essential role, largely determining the sensitivity and spatial resolution of these techniques. It would be ideal if human tissues were as transparent as those of a jelly fish or X-ray fish. However, due to the inhomogeneities of the refractive index among different cells and organelles in tissue [11], photons are multiply scattered along their respective optical paths. As a result, an incident optical beam spreads out and becomes diffusive very quickly as it travels deep into tissue (typically \(\sim\)1 mm beneath human skin for visible and near-infrared light). Beyond this so-called optical diffusion limit, optical focus can hardly be formed by using lenses or objects; instead, if the coherence length of light is sufficiently long, optical speckle patterns are developed within or behind a thick biological tissue sample. Moreover, because of the strong scattering, the in situ optical fluence also decays rapidly. These two folds pose a fundamental limitation to the aforementioned high-resolution techniques to be viable only within superficial depths (e.g., less than 100-200 micrometers). Recently, photoacoustic (PA) imaging has been introduced to tackle the diffusion-induced resolution barrier in deep tissue [11; 12; 13; 14; 15]. This technique is based on the so-called PA effect [16; 17; 18], which mainly consists of four steps: first, tissue molecules in the region of interest (ROI) within the sample absorb pulsed light (no matter if it is in ballistic or diffusive form); the absorbed optical energy is converted to heat, causing local the temperature to increases; ultrasonic waves are generated due to thermal expansion; finally, these ultrasonic waves (typically around the order of MHz if the exiting light pulse width is several nanoseconds) can be detected by one or a series of ultrasound transducers mounted outside the sample. Since ultrasonic waves are scattered almost 1000 times less than light is in tissue (unless otherwise mentioned, we always mean soft tissue in the context), they can be used to reconstruct an image of the absorbing ROI with an acoustic resolution [19; 20; 21; 22; 23; 24; 25; 26; 27]. As the generation of such PA signals does not distinguish ballistic or diffusive exciting light formations, it allows for a penetration depth far exceeding the optical diffusion limit. On the other hand, if only shallow tissue depths (e.g., within one transport mean free path) are sought, where ballistic or quasi-ballistic photons dominate, the optical spot size can be much smaller than that of the ultrasonic focus, the resolution of PA imaging is determined by the optical--but not the acoustic--focal size. In this scenario, the technique is often referred to as optical-resolution photoacoustic imaging or microscopy [28; 29]. As a hybrid modality, PA imaging takes advantage of the optical contrast and acoustic or optical spatial resolution at moderate tissue depth. Therefore, it has recently seen many fundamental and preclinical trials for the imaging or diagnosis of blood vessels [30], brain [31], breast tumor [32], sentinel lymph node (SLN) [33], oxygen saturation [34], sub-cellular structures [14], and small-animal whole body [35], among many other explorations.