LUS Results and Human Imaging

The LUS imaging results for porcine and human tissue demonstrate the capability of LUS to remotely and deeply image biological tissue at safe optical exposure levels. The tissue boundaries detected by LUS are in agreement with the tissue boundaries detected by conventional ultrasound. Components of LUS technology have been explored in prior research, but a full demonstration of LUS on a human subject has not previously been presented. The human LUS imaging presented here (Fig. 4c) is the first instance of LUS imaging on a human subject in vivo, and volumetric LUS imaging (Fig. 2c) demonstrates LUS capabilities for 3D imaging. By restricting optical penetration and selecting optical wavelengths that maximize both the generation and detection of acoustic waves, we can demonstrate LUS imaging on tissue-mimicking phantoms, excised porcine tissue, and human subjects without any surface enhancements for the optical source or detector performance. Validation of the LUS results on human subjects is a significant step toward proving the clinical viability of LUS and motivates further LUS research and development.

LUS imaging in porcine tissue (Fig. 3c) shows a clear and coherent boundary layer in the first 1.5 cm, but deeper layers present significant artifacts near the boundary. The same effect is visible in the human LUS image (Fig. 4c) for deeper layers. These depth-dependent artifacts are likely due to the lack of elevation focusing in LUS. While a conventional ultrasound probe is elevation focused to minimize out-of-plane reflections, optical sources and receivers are unfocused and become sensitive to more out-of-plane reflection at deeper imaging depths. More artifacts are visible in the human LUS image (Fig. 4c) than in the porcine LUS image (Fig. 3c) and can be attributed to patient motion and variations in optical backscatter from the skin. Overall, less acoustic speckle is present in the LUS images due to the lower imaging frequency, bandwidth, and lateral resolution of LUS than those of conventional ultrasound. Currently, LUS performance is limited by the acoustic frequency generated by the optical source and the detector sensitivity. The frequency of LUS sources is determined by the impinging optical wavelength and tissue absorption, with bandwidth limits due to the acoustic attenuation and tissue surface roughness; the expected source bandwidth for LUS at 1540 nm is ~1.5 MHz45. For optical detection, noncontact interferometric methods are limited by the optical backscatter from the tissue surface. Optical detectors have more than sufficient bandwidth for ultrasound imaging but are generally reliant on ideal stationary reflective surfaces for measurement, without consideration of human safety35. Based on our human subject experiments, there is significant variation in the optical backscatter, as the detector was used to scan the skin tissue. Optical reflectance on human subjects has been reported for optical wavelengths between 250 and 2500 nm, but the measurement of backscatter specularity is lacking50,51. Full skin reflection characteristics must be investigated to design specialized optical detectors for clinical LUS. For the LUS system reported here, the 1550 nm LDV was specifically selected to maximize permissible optical backscatter from the skin while maintaining safety. Adaptive focusing could further reduce variability associated with skin variations and patient motion, but feedback control of the optical output based on the optical backscatter will be necessary.

At this point, the LUS images presented here are comparable to images presented at the incipient stages of medical ultrasound imaging decades ago1. Quantitative image comparison of the current LUS technology against modern medical ultrasound is still premature. The conventional ultrasound images presented here use a 9 MHz center frequency transducer with multielement beamforming and are of expectedly higher image quality and resolution than those of the LUS images. Nevertheless, similar structures and dimensions are consistently observed in both modalities. While further work remains prior to commercialization and clinical use, the core enabling technologies of LUS are available. Discussed in later sections, recent research advances in laser technology, silicon photonics, and hydrogels could accelerate future LUS development.

Optical system design

From prior literature, LUS sources can be approximated by an equivalent disk transducer below the tissue surface, with the geometry determined by the parameters of the impinging light and absorption characteristics of the irradiated tissue22,23. With proper thermal and mechanical stress confinement, the equivalent disk is prescribed, in diameter, by the impinging optical beam diameter, and in thickness, by the optical penetration depth (inverse of the tissue optical absorption coefficient at the impinging optical wavelength)22. Thus, the central frequency of the converted acoustic wave and the conversion efficiency are directly determined by the optical absorption coefficient22. For practical applications, the beam diameter should be larger than the penetration depth such that the dominant propagation direction of the converted acoustic wave is directed into the tissue. For maximum conversion efficiency, the optical absorption coefficient should equal the acoustic wavenumber of the desired acoustic frequency to be generated22. For biological tissue, the stress confinement conditions for LUS are satisfied by nanosecond pulsed infrared sources, restricting optical conversion at the tissue surface using high optical absorption regions of water45. To generate acoustic frequencies relevant for clinical ultrasound (1–10 MHz), ideal optical wavelengths for LUS source generation are ~1500 and 2000 nm, corresponding to peaks in the water absorption spectrum. At 1500 nm, the optical absorption water is ~40 cm−1, which corresponds to a converted acoustic source near 1 MHz. Prior work reports that both 1500 nm and 2000 nm pulsed sources generate acoustic waves with bandwidths up to 1.5 MHz45. However, the MPE limit for nanosecond sources at 2000 nm is an order of magnitude lower than that at 1500 nm, 0.1 J/cm2 vs. 1 J/cm2, respectively46. Thus, the 1540 nm optical source was selected to maximize the converted acoustic source amplitude while remaining within the safety limits. Following the disk transducer model, the optical beam diameter dictates the beam profile of the converted acoustic source. For the 2 mm diameter optical source on tissue with an expected frequency of ~1.5 MHz, the acoustic beam width is approximately 60 degrees. A narrower beam spread can be generated using larger optical beam diameters, but significantly higher optical energy is required to maintain the overall fluence.

For optical detection, optical backscatter from the tissue surface must be maximized. Prior measurements of optical reflectance from human skin report that wavelengths between 500 nm and 1200 nm offer the highest optical reflection factor46,50,51. However, the safety MPE limits must be considered again. For a 100 µs measurement duration (corresponding to an imaging depth of 7.5 cm in tissue), assuming similar photodetector responsivity across the spectrum, if a tissue surface is irradiated at the maximum MPE limits for every spectral region, the maximum quantity of reflected light to a detector is still in the far infrared, between 1500 and 1800 nm. Thus, the selected 1550 nm LDV balances detection sensitivity with subject safety. Considering subject safety and acoustic performance, ~1500 nm is the optimal spectral region for both LUS transmission and detection and is able to leverage the commercial abundance of 1550 nm optical components from the telecommunication industry45.

LUS performance is directly linked to the intrinsic tissue properties of human skin. The absorption and reflection parameters dictate LUS transmission and detection, respectively. The tissue surface roughness limits the LUS transmit bandwidth and may contribute to variation and specularity in optical reflection, which degrades the optical detection sensitivity. Skin tissue inhomogeneity from pigmentation and melanin may further affect the LUS performance and warrant further investigation. Averaging multiple traces improves the overall signal-to-noise ratio, but the method is limited by MPE limits and increases the overall imaging time; poor or a lack of optical backscatter from the tissue surface cannot be resolved by averaging. Surface treatment, adaptive focusing, or multichannel optics may be necessary for clinical LUS systems to mitigate tissue variability. Furthermore, unlike fixed-geometry ultrasound arrays, the equivalent optical LUS array conforms to the local tissue geometry. The array geometry for the LUS images presented can be approximated as planar, but the imaging of large tissue regions with significant local curvature requires active localization of the sources and detectors to accurately reconstruct an image.

Enabling LUS technologies

Future LUS development should focus on component improvements as well as surface treatments to enhance multiple facets of LUS. An amplitude-modulated optical source can excite higher acoustic bands to improve the image resolution22. Implementing a fast amplitude-modulated optical source could improve the LUS imaging depth, bandwidth, and resolution; in addition, conventional ultrasound techniques used to amplify the acoustic signal-to-noise ratio, such as pulse compression, matched filtering, and source encoding, can be leveraged. Currently, both LUS and PA are limited by the existing laser technology16,17,18,41,52. Beyond amplitude modulation, the parallelization of optical sources and receivers to enable optical transmit and receive beamforming techniques will be a critical turning point, similar to how piezoelectric arrays enabled and revolutionized clinical ultrasound imaging. Since the spatial locations of LUS sources are restricted to the tissue surface and can be known a priori, transmit beamforming is possible only in LUS and not in penetrating PA. The adaptation of computer vision techniques such as 3D imaging and tracking technologies could make large-volume LUS imaging and optical beamforming feasible. Optical spot tracking is also required if patient motion is significant during data acquisition. With sufficiently fast data acquisition and coverage, LUS systems could behave similar to a body volume camera, capable of simultaneous 3D imaging of both the external and internal tissue geometries. The silicon photonics industry, fueled by sensors required for autonomous car operations, has developed chip-scale, solid-state, steerable laser technology applicable for both PA and LUS imaging53,54,55,56,57. Since the ideal operating range for LUS is near 1500 nm, communication and silicon photonic innovations may directly impact LUS system development.

As discussed previously, the LUS system performance is subject to the optical and acoustic properties of the tissue. A minor surface treatment layer could bypass the tissue limitations. Radio frequency (RF) coils and contrast agents are commonly used for MRI and CT, respectively; designing a surface treatment layer for LUS is not inconceivable. Enhanced optical reflection is commonly achieved with retroreflective dust or tapes39,42. However, retroreflective dust is a respiratory irritant, and tape or dust on the tissue surface impedes the LUS source. Gel or gel pads embedded with retroreflectors could be designed to enhance optical detection without interference to the optical source; the bulk water content of the gel can preserve the LUS source characteristics in the IR spectrum, while embedded reflectors can enhance the optical backscatter. Furthermore, source bandwidth limitations due to the tissue roughness can be bypassed by generating the LUS source in the gel layer. Gel or gel pads routinely used for ultrasound imaging can be augmented for LUS imaging. Flexible hydrogels could also be designed for LUS; the bulk water content of hydrogels mixed with retroreflective particles can conform to the tissue surface and enhance the LUS performance58. The treatment layer could expand the LUS source bandwidth, improve the optical detection sensitivity, and permit higher optical exposure limits. However, irradiance exceeding the MPE limits on the treatment layer may mandate additional safety measures such as enclosure or eye protection. New optical designs should be evaluated to accommodate the use of surface treatments in LUS imaging.

Outlook

Based on these encouraging results, LUS inspires confidence for further research and development. Human LUS images verify the feasibility of LUS for in vivo anatomical imaging without compromising patient safety. LUS is sensitive to both hard and soft anatomical features, similar to conventional ultrasound, but is fully noncontact. The real-time remote sensing of biological tissue would find broad applicability in nonintrusive patient monitoring, contact sensitive imaging (elastography and musculoskeletal disease tracking), and intraoperative applications. Current embodiments of LUS for research are generally single-point transmission and detection due to the high cost and complexity. Clinical iterations of LUS will require multipoint optical transmission and detection to amplify the acoustic source amplitudes and reduce the data acquisition time. Analogous challenges existed in the nascent stages of conventional ultrasound imaging. Scaling from mechanically scanned single-element systems to highly parallelized real-time clinical imagers took decades of research and development1. A similar pathway is likely for LUS – parallelization of single- to multipoint laser technology. The initial human results are encouraging; rapid advances in related industrial sectors will establish the technologies necessary to enable clinical implementation of LUS.