WO2025129187A1 - System and method for extended and/or continuous ultrasound elastography - Google Patents

Abstract

A method for manufacturing a wearable ultrasound probe is provided. The method includes forming a plurality of slits in a layer of piezoelectric material. Each slit has a kerf size less than 40 μm and a center-to-center separation between 150 μm and 250 μm. The slits define edges of a plurality of ultrasound transducer elements that are formed in the piezoelectric material. The ultrasound transducer elements are configured to produce an acoustic radiation force impulse (ARFI) in a tissue. The method further includes adhering the layer of piezoelectric material to a matching layer and implanting a bio-adhesive layer onto the matching layer. The bio-adhesive layer is configured to adhere to skin of a subject.

Description

SYSTEM AND METHOD FOR EXTENDED AND/OR CONTINUOUS ULTRASOUND

ELASTOGRAPHY

CROSS REFERENCE TO RELATED APPLICATIONS

[0001 ) This application is based on, claims priority to, and incorporates herein by reference for all purposes, U.S. Provisional Patent Application No. 63/610,805 filed on December 15, 2023.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT [0002] This invention was made with government support under EY032229, HL153857, HL167947, and EB032958 awarded by the National Institutes of Health. The government has certain rights in the invention.

BACKGROUND

[0003] Fulminant hepatic failure, also known as acute liver failure (ALF), is defined as a severe liver injury resulting in the onset of hepatic encephalopathy within 8 weeks of the initial symptoms in patients without underlying liver disease. ALF has a high mortality rate of approximately 80% due to massive short-term cell death. This condition can occur as a result of various etiologies such as viral hepatitis (hepatitis A and E and hepatitis B), neoplastic infiltration, heart failure, my cot oxi cosis, drug toxicity, or complications of liver transplantation. The survival rate of patients with ALF after one month is merely 23% if appropriate procedures such as liver transplantation or intensive care medicine are not immediately taken. Even after liver transplantation, post-operative complications may still lead to acute liver graft dysfunction and high premature mortality. Therefore, prompt prognostic evaluation plays an important role in performing timely intensive care treatment of ALF, allowing graft salvage, and managing postoperative complications in the intensive care unit (ICU).

[0004] The gold standard for prognosis models of ALF is mainly based on the histological examination of liver tissue biopsies, laparoscopy for liver surface evaluation, and laboratory markers. However, these methods are time-consuming, have a certain level of risks due to invasive procedures, and could be impractical for critically ill patients. Several studies have reported that liver stiffness is a promising prognosis marker for early detection of ALF. Critically ill patients with ALF in the ICU have shown a significantly increased liver stiffness (Young’s modulus of approximately 27 kPa) compared to healthy controls (Young’s modulus of approximately 3.8-6 kPa). Ultrasound shear wave elastography (SWE) is a fast and noninvasive technique used to evaluate the stiffness of internal organs. However, ultrasound elastography requires clinicians to use hand-hold ultrasound probes on patients during examinations, making it unable to continuously monitor changes in liver stiffness over the course of rapidly changing diseases.

|0005| The concept of wearable ultrasound can be traced back as early as 2005 when a pair of ultrasound transducers was implanted into the fracture region of mature sheep to monitor and accelerate the healing process of fractured long bones. Over the past decade, with the rapid growth of soft materials, wearable ultrasound devices have been extensively developed for imaging internal organs and evaluating various physiological features in biomedical applications such as measuring central blood pressure, assessing deep-tissue hemodynamics, monitoring cardiac functions, and continuously imaging diverse organs. However, these methods are primarily focused on imaging structures of biological tissues, not available for elastography measurements (Table 1, FIG. 11).

[0006] A recent report introduced a wearable device based on a stretchable ultrasound probe for measuring muscle stiffness using a compression-based excitation method. However, this method has limitations when it comes to the continuous measurement of stiffness of internal organs such as liver (Table 1, FIG. 11). The compression-based approach requires an additional tool, such as a finger or an actuator, to induce mechanical excitation; it is impractical to repeatedly compress one location of the skin over the course of the diseases. Additionally, compressing the skin may be unable to mechanically excite internal organs such as the liver and thus unsuitable for measuring their moduli. Furthermore, mechanical compression can distort the acoustic beam and stretchable ultrasound probe, which dramatically compromises the performances of ultrasound elastography (Table 1, FIG. 11).

[0007] Accordingly, there is an ongoing need for improved methods and systems for providing continuous and accurate monitoring of patents with serious and rapidly-changing conditions, such as ALF.

SUMMARY

[0008] In some aspects, a method for manufacturing a wearable ultrasound probe is provided.

The method includes forming a plurality of slits in a layer of piezoelectric material. Each slit has a kerf size less than 40 pm and a center-to-center separation between 150 pm and 250 pm. The slits define edges of a plurality of ultrasound transducer elements that are formed in the piezoelectric material. The ultrasound transducer elements are configured to produce an acoustic radiation force impulse (ARFI) in a tissue. The method further includes adhering the layer of piezoelectric material to a matching layer and implanting a bio-adhesive layer onto the matching layer. The bio-adhesive layer is configured to adhere to skin of a subject.

[0009] In other aspects, a wearable ultrasound probe, which is configured to measure shear wave elastography, is provided. The probe includes an ultrasound array that has a plurality of transducer elements. Each transducer element includes a piezoelectric material. The probe further includes a bio-adhesive layer that is coupled to the plurality of transducer elements. The bio-adhesive layer is configured to adhere to skin of a subject to position the ultrasound array to acquire shear wave elastography data from the subject. One or more of the plurality of transducer elements is configured to produce an ARFI that excites a shear wave in a tissue of the subject. One or more of the plurality of transducer elements is configured to measure shear wave elastography data generated in response to the shear wave.

[0010] In still other aspects, a method for measuring stiffness of a tissue is provided. The method includes placing a wearable ultrasound probe on skin of a subject. The wearable ultrasound probe includes a bio-adhesive that is configured to adhere the probe onto the skin of the subject. The probe also includes a plurality of ultrasound transducers that are configured to produce an ARFI in a subcutaneous tissue of the subject. The method further includes using one or more of the plurality of ultrasound transducers to produce an ARFI to generate a shear wave in a tissue of the subject. The method also includes using one or more of the plurality of ultrasound transducers to measure particle displacement data that characterize particle displacement in the tissue generated in response to the shear wave. The method also includes using a processor to determine a stiffness of the tissue based on the particle displacement data.

[0011] The foregoing summary is illustrative only and is not intended to be in any way limiting. In addition to the illustrative aspects, embodiments, and features described above, further aspects, embodiments, and features will become apparent by reference to the drawings and the following detailed description. BRIEF DESCRIPTION OF THE DRAWINGS

[0012] For a further understanding of the nature, objects, and advantages of the present disclosure, reference should be made to the following detailed description, read in conjunction with the following drawings, wherein like reference numerals denote like elements and wherein:

[0013] FIG. 1A shows a cross-section of a wearable bio-adhesive ultrasound elastography probe.

10014] FIG. IB provides a flowchart setting forth steps of a method of manufacturing a wearable bio-adhesive ultrasound elastography probe.

[0015] FIG. 1C provides a flowchart setting forth steps of a method of performing ultrasound elastography using a wearable bio-adhesive ultrasound elastography probe.

[001.6] FIG. 2A is a schematic illustration of a BAUS-E system in accordance with the present disclosure, demonstrating an ability to provide an ARFI push to generate shear waves in biological tissues and to detect particle displacements induced by an ARFI push in biological tissues.

[0017] FIG. 2B is a schematic illustration of the BAUS-E components of FIG. 2A, including the backing layer, flexible printed circuit board (f-PCB), piezoelectric composite, matching layer, and bio-adhesive couplant.

[0018] FIG. 2C is a schematic illustration of a BAUS-E device adhered to the skin, demonstrating its ability to withstand a high pulling force and maintain robust adhesion.

[0019] FIG. 3A is a graph created in one non-limiting example showing the pulse-echo performance and -6dB fractional bandwidth measurement of 50^th element of the BAUS-E. The HF and LF at -6 dB were 8.34 MHz and 6.06 MHz, respectively. The center frequency of the 50^th element of the BAUS-E was 7.2 MHz.

[0020] FIG. 3B is a graph created in one non-limiting example that illustrates the simulated pulse-echo performance and bandwidth by PiezoelectricCAD simulation results. The HF and LF at -6 dB were 10.30 MHz and 5.90 MHz, respectively, and the center frequency in the simulation was 8.1 MHz.

[0021 ] FIG. 3C is a graph created in one non-limiting example shows the center frequency of all 128 elements of the BAUS-E and the mean center frequency of the BAUS-E was 7.52 ± 0.26 MHz, which demonstrates our outstanding transducer fabrication technique to make the center frequency of the BAUS-E that is close to the PeizoCAD simulation result. HF: higher frequency; LF: lower frequency.

[0022] FIG. 4 A is a graph that shows the impedance and phase measurement of 50^th element of the BAUS-E. The /a and f- happened at 8.60 MHz and 6.90 MHz, respectively. The impedance Im of the 50^th element of the BAUS-E was approximately located as 165 Q.

[0023] FIG 4B is a graph that illustrates the simulated impedance and phase spectrum using the PiezoelectricCAD simulation. The fa an fr in the simulation appeared at 8.50 MHz and 6.70 MHz, respectively. The simulated impedance Im of the 50^th element was approximately 164 .

[0024] FIG. 4C is a graph that shows the impedance of all 128 elements of the BAUS-E and the mean impedance of the BAUS-E was 164.73 ± 2.76 , which is a good agreement with the simulation result, fa. anti-resonance frequency; : resonance frequency; Im transducer impedance.

[0025] FIG. 5 A graph that shows the peak-peak sensitivity of the BAUS-E wearable device (red circle) and the commercial transducer (Verasonics system equipped with L7-4 probe) (blue circle). The sensitivity of the BAUS-E device and the commercial transducer were -3.18 ± 1.52 dB and -4.15 ± 2.36 dB, respectively.

[0026] FIG. 5B is a graph showing the fractional bandwidth of the BAUS-E and compared with the commercial transducer. The fractional bandwidth of the BAUS-E and the commercial transducer were 37.45 ± 4.86% and 44.35 ± 3.05%.

[0027] FIG. 6A is an image from a lateral and axial resolution test of the BAUS-E showing the lateral resolution of the BAUS-E.

[0028] FIG. 6B is a graph of axial resolution of the BAUS-E. The axial resolution at -6B was approximately 0.36 mm according to the tungsten wire target at approximate 13 mm.

[0029] FIG. 6C is a graph of lateral resolution The lateral resolution at -6 dB was approximately 0.13 mm.

[0030] FIG. 7 is a set of correlated graphs of numerical simulation and their dispersion analysis of the phantom study using the commercial transducer (L7-4) and the BAUS-E w/wo bio-adhesive couplant.

(0031] FIG. 8 is a set of correlated graphs from systematic phantom studies to evaluate the accuracy of shear waves.

[0032] FIG. 9 is a set of correlated graphs and images. (0033] FIG. 10A is a schematic illustration depicting the procedure for evaluating elasticity changes in rats with ALF. D-Galactosamine was injected into nine rats via intraperitoneal injection to induce ALF. The elasticity of the rat liver was continuously monitored for 48 hours with 6-hour intervals using BAUS-E to obtain SWEs. At each 6-hour interval, one rat was randomly selected to sacrifice to obtain liver specimens for pathological staining, which corresponded to the timeline of the elasticity measurements. Regression analysis was performed to explore the relationship between pathological results and the shear modulus of rats with ALF.

[0034] FIG. 1 OB is a set of correlated spatiotemporal maps illustrating shear wave velocities measured at 0, 12, 30, and 48 hours as examples.

[0035] FIG. 10C is a graph of shear waves velocities were observed from 0 to 48 hours using the BAUS-E.

[0036] FIG. 10D is a graph showing a trend of Young's modulus changes in relation to the severity of rats with ALF over 48 hours. IF: immunofluorescence. *p < 0.05, **p < 0.01, and ***p < 0.001.

[0037] FIG. 11. Table 1 : The comparison of performances of the BAUS-E device with existing wearable ultrasound devices.

[0038] FIG. 12. Table 2: The fabrication parameters of the BAUS-E.

[0039] FIG. 13 is a block diagram of a non-limiting example ultrasound system that can implement the methods described in the present disclosure.

|0040] FIG. 14 is a profile view of a non-limiting example wearable ultrasound system.

[0041] FIG. 15 is a cross-section of a non-limiting example wearable ultrasound system.

DETAILED DESCRIPTION

[0042] Reference will now be made in detail to presently preferred embodiments and methods of the present invention, which constitute the best modes of practicing the invention presently known to the inventors. The Figures are not necessarily to scale. However, it is to be understood that the disclosed embodiments are merely exemplary of the invention that may be embodied in various and alternative forms. Therefore, specific details disclosed herein are not to be interpreted as limiting, but merely as a representative basis for any aspect of the invention and/or as a representative basis for teaching one skilled in the art to variously employ the present invention. (0043] It is also to be understood that this invention is not limited to the specific embodiments and methods described below, as specific components and/or conditions may, of course, vary. Furthermore, the terminology used herein is used only for the purpose of describing particular embodiments of the present invention and is not intended to be limiting in any way.

(0044] It must also be noted that, as used in the specification and the appended claims, the singular form "a," "an," and "the" comprise plural referents unless the context clearly indicates otherwise. For example, reference to a component in the singular is intended to comprise a plurality of components.

[0045] The term “comprising” is synonymous with “including,” “having,” “containing,” or “characterized by.” These terms are inclusive and open-ended and do not exclude additional, unrecited elements or method steps.

|0046( It should also be appreciated that integer ranges explicitly include all intervening integers. For example, the integer range 1-10 explicitly includes 1, 2, 3, 4, 5, 6, 7, 8, 9, and 10. Similarly, the range 1 to 100 includes 1, 2, 3, 4. . . . 97, 98, 99, 100. Similarly, when any range is called for, intervening numbers that are increments of the difference between the upper limit and the lower limit divided by 10 can be taken as alternative upper or lower limits. For example, if the range is 1.1. to 2.1 the following numbers 1.2, 1.3, 1.4, 1.5, 1.6, 1.7, 1.8, 1.9, and 2.0 can be selected as lower or upper limits.

(0047] When referring to a numerical quantity, in a refinement, the term “less than” includes a lower non-included limit that is 5 percent of the number indicated after “less than.” A lower nonincludes limit means that the numerical quantity being described is greater than the value indicated as a lower non-included limited. For example, “less than 20” includes a lower non-included limit of 1 in a refinement. Therefore, this refinement of “less than 20” includes a range between 1 and 20. In another refinement, the term “less than” includes a lower non-included limit that is, in increasing order of preference, 20 percent, 10 percent, 5 percent, 1 percent, or 0 percent of the number indicated after “less than.”

(0048) With respect to electrical devices, the term “connected to” means that the electrical components referred to as connected to are in electrical communication. In a refinement, “connected to” means that the electrical components referred to as connected to are directly wired to each other. In another refinement, “connected to” means that the electrical components communicate wirelessly or by a combination of wired and wirelessly connected components. In another refinement, “connected to” means that one or more additional electrical components are interposed between the electrical components referred to as connected to with an electrical signal from an originating component being processed (e.g., filtered, amplified, modulated, rectified, attenuated, summed, subtracted, etc.) before being received to the component connected thereto.

[0049] The term “electrical communication” means that an electrical signal is either directly or indirectly sent from an originating electronic device to a receiving electrical device. Indirect electrical communication can involve processing of the electrical signal, including but not limited to, filtering of the signal, amplification of the signal, rectification of the signal, modulation of the signal, attenuation of the signal, adding of the signal with another signal, subtracting the signal from another signal, subtracting another signal from the signal, and the like. Electrical communication can be accomplished with wired components, wirelessly connected components, or a combination thereof.

[0050] The term “one or more” means “at least one” and the term “at least one” means “one or more.” The terms “one or more” and “at least one” include “plurality” as a subset.

[0051] The terms “substantially,” “generally,” or “about” may be used herein to describe disclosed or claimed embodiments. The term “substantially” may modify a value or relative characteristic disclosed or claimed in the present disclosure. In such instances, “substantially” may signify that the value or relative characteristic it modifies is within ± 0%, 0.1%, 0.5%, 1%, 2%, 3%, 4%, 5% or 10% of the value or relative characteristic.

|0052] The term “electrical signal” refers to the electrical output from an electronic device or the electrical input to an electronic device. The electrical signal is characterized by voltage and/or current. The electrical signal can be stationary with respect to time (e.g., a DC signal) or it can vary with respect to time.

[0053] The term “electronic component” refers is any physical entity in an electronic device or system used to affect electron states, electron flow, or the electric fields associated with the electrons. Examples of electronic components include, but are not limited to, capacitors, inductors, resistors, thyristors, diodes, transistors, etc. Electronic components can be passive or active.

[0054] The term “electronic device” or “system” refers to a physical entity formed from one or more electronic components to perform a predetermined function on an electrical signal. (0055] The processes, methods, or algorithms disclosed herein can be deliverable to/implemented by a processing device, controller, or computer, which can include any existing programmable electronic control unit or dedicated electronic control unit. Similarly, the processes, methods, or algorithms can be stored as data and instructions executable by a controller or computer in many forms including, but not limited to, information permanently stored on non-writable storage media such as ROM devices and information alterably stored on writeable storage media such as floppy disks, magnetic tapes, CDs, RAM devices, and other magnetic and optical media. The processes, methods, or algorithms can also be implemented in a software executable object. Alternatively, the processes, methods, or algorithms can be embodied in whole or in part using suitable hardware components, such as Application Specific Integrated Circuits (ASICs), Field-Programmable Gate Arrays (FPGAs), state machines, controllers or other hardware components or devices, or a combination of hardware, software and firmware components.

[0056] Abbreviations:

[0057] “ARFI” means acoustic radiation force impulse.

[0058] “BAUS-E” means wearable bio-adhesive ultrasound shear wave elastography.

[0059] There is a need for ultrasound elastography technology that can provide continuous measurements of internal organ stiffness over several hours. The present disclosure provides a wearable ultrasound elastography technology that can continuously monitor the evolution of liver stiffness at the early stage of ALF, which can aid the prognosis evaluation of ALF. The wearable ultrasound described herein is capable of providing acoustic radiation force impulses (ARFI) that can generate shear waves in biological tissues for elastography.

[0060] In order to produce shear waves that are excited by ARFI, an ultrasound device must use a very high voltage. Existing wearable ultrasound devices, such as piezoelectric micromachined ultrasonic transducers (PMUT) and capacitive micromachined ultrasonic transducer (CMUT) are not capable of such high voltages without burning out the electronics of the system. The high voltage requirement is further complicated by the desire to have a small device that is easily and comfortably wearable, in contrast to existing diagnostic systems that have apertures 40 mm or larger that provide around 90 V. The systems and methods of the present disclosure provide a solution to such challenges using a lead zirconate titanate (PZT)-based array that can provide a high voltage with a very small aperture size (e.g., ~25 mm). As a non-limiting example, the array can be preferably operated with a voltage between 30-50 V without damaging the device. As other examples, the array may be configured to be operated with voltages of 10-20 V, 20-30 V, 30-40 V, 40-50 V, or 50-60 V. The PZT material can be diced to provide a multi-channel array that improves imaging quality. Such ultrasound probe is capable of generating radiation force that produce shear waves in biological tissues, that can be used for ultrasound elastography to measure tissue stiffness.

[0061] In at least one aspect, a wearable bio-adhesive ultrasound shear wave elastography probe is provided. The probe includes an ultrasound piezoelectric material layer having a plurality of transducer elements that can form an ultrasound beam. The transducer elements can be arranged in a linear array. The plurality of transducer elements configured to allow electronic focusing and steering of the ultrasound beam, and to provide acoustic radiation force impulses (ARFI). A portion of the transducer elements are configurable to generate acoustic radiation force impulses as an excitation source to produce shear waves for elastography measurements. The probe also includes a bio-adhesive hydrogel layer coupled to the ultrasound piezoelectric material layer. Advantageously, the bioadhesive hydrogel layer is configured to adhere the wearable ultrasound shear wave elastography probe to a subject’s skin for continuous and wearable ultrasound elastography for a predetermined time period (e.g., several hours).

[0062] The wearable bio-adhesive ultrasound elastography (BAUS-E) is capable of continuously measuring liver stiffness over a 48-hour period is provided. This enables the evaluation of the rapid changes in liver stiffness associated with ALF for prognostic purposes. In some implementations, the BAUS-E incorporates a thin transducer array with 128 channels that is merely 24 mm in azimuth direction. The thin transducer array can generate acoustic radiation force impulses (ARFI) as an excitation source to produce shear waves for elastography measurements. The thin array is further adhered to the skin via a bio-adhesive tough hydrogel couplant for continuous and wearable ultrasound elastography for over 48 hours. This bio-adhesive hydrogel-elastomer hybrid couplant can maintain high water content (>90 wt%), high interfacial toughness (>500 Jm’²), and low acoustic attenuation (i.e., similar to commercial ultrasound gels) for continuous and wearable ultrasound elastography. The small aperture size of the BAUS-E device with the soft bio-adhesive couplant provide a comfortable wearability. The performances of an example implementation of the BAUS-E is systematically characterized herein, including transducer measurements, phantom studies, and numerical simulations. The performance of the BAUS-E was further validated in comparison with a commercial ultrasound transducer. Additionally, the BAUS-E was shown to continuously measure liver stiffness changes in rats with ALF induced by D-Galactosamine over 48 hours. The increase in liver stiffness measured by BAUS-E was significantly correlated with ALF indicated by histopathological staining. The BAUS-E could find extensive applications in clinics and ICU for noninvasive and continuous monitoring of the moduli of internal organs over the course of rapidly changing diseases, such as ALF, for their early detection.

[0063] Referring to FIG. 1A, a cross-section of an ultrasound probe 10 is provided. The ultrasound probe 10 can be wearable in order to continuously measure (e.g., once every 1 minute, once every 10 minutes, once every hour, once every day) shear wave elastography of a subject or patient. The probe may be worn by the user for an extended period of time (e.g., greater than 4 hours, up to 12 hours, up to 24 hours, up to 48 hours, up to one week) to longitudinally measure elastography. The described probe can advantageously produce acoustic radiation force impulses (ARFI) as an excitation source to produce shear waves in the tissue that can be used to assess tissue characteristics, such as liver stiffness.

[0064] The probe 10 contains a piezoelectric material layer 12 that includes a plurality of transducer elements 14 arranged in an ultrasound array (e.g., a linear array). As non-limiting examples, the array may include 64, 128, 256, or 512 elements. Each element may be configured as a transmit element, a receive element, or a transmit/receive element. As one non-limiting example, a subset of the elements (e.g., 32 elements) may be configured as transmit elements. The transmit elements can produce an ultrasonic beam that can be focused and steered to produce the ARFI and excites a shear wave in the tissue. The number of transmit elements can be tuned based on the desired penetration depth. For example, the number of transmit elements may be increased to reach deeper anatomical structures or decreased to provide a stronger shear wave at the tissue of interest. All or a subset of the elements (e.g., 128) may be configured as receive elements to measure the particle displacement in the tissue in response to the shear waves produced by the ARFI. This particle displacement can be used to characterize tissue stiffness.

[0065] The piezoelectric material 12 may include lead zirconate titanate (PZT). In some configurations, the piezoelectric material may include a 1-3 composite PZT material. The piezoelectric material 12 may also include electrodes on one or both sides of the piezoelectric layer 12. (0066] Transducer elements may be formed in the piezoelectric material layer 12 to produce an electrode separation that can be described by a pitch 32 and kerf 34 size. In some implementations, each transducer element 14 may be generally configured as a rectangular slab (i.e., a rectangular parallelepiped). The electrode separation can be achieved by scratch-dicing the piezoelectric layer 12 using a dicing saw. In some implementations the dicing saw has a thickness no greater than that of the kerf size 34. In some configurations, the space between each transducer element 14 may be filled with epoxy.

[0067] As one non-limiting example, each transducer element may have a pitch size 32 defined based on the shear-wave wavelength ( ) of the target media, which may be approximated as X of water (Xwater). Thus, the pitch size may be approximately 200 pm. In some configurations, the kerf size 34 or element separation may be defined based on the criterion JX/4/c < 40 pm, where Vs represents the shear velocity within the epoxy-filled kerf and f_c is the center frequency of the array. In one nonlimiting example, the kerf size may be 20 pm.

[0068] The piezoelectric material 12 can also be described by a thickness 36 (e.g., along z) and a width (e.g., along y, which may be referred to as the elevation direction). It may be advantageous to constrain the width-to-thickness ratio in order to generate higher electromechanical coupling and enable pure thickness mode resonance, which could provide better performance. As a non-limiting example, the width-to-thickness ratio can be constrained as <0.7. A higher width-to-thickness ratio (e.g., >0.7) could induce different resonance vibration modes that interfere with the fundamental resonance mode, which could reduce the bandwidth and signal-to-noise ratio (SNR). As one nonlimiting example, the transducer may have a width of 80 pm and a thickness of 140 pm to provide a width-to-thickness ratio of 0.57. As other non-limiting examples, the transducer may have a width of 50-100 pm or 70-90 pm and a thickness of 100-200 pm or 120-160 pm.

[0069] The piezoelectric material layer 12 that forms an ultrasound array may also include an electrode layer on one or both sides of the piezoelectric material layer. Such electrode may include a conductive material that transfers the mechanical stress applied to the piezoelectric material to an electrical signal or voltage and vice versa. The electrode can be coupled to ultrasound circuitry in a circuit layer 18. The circuit layer 18 may include a printed circuit board (PCB) or a flexible PCB (f- PCB) that can transmit the electrical signals to and from a processor or data storage device. In this way, the f-PCB can be disposed over the piezoelectric layer 12. In some configurations, there may be an additional circuit layer on the bottom size of the piezoelectric layer 12 between the transducer elements 14 and the matching layer 24.

[0070] The probe 10 also includes a bio-adhesive layer 16 coupled to the ultrasound array. The bio-adhesive layer 16 includes a biocompatible adhesive, such as a hydrogel, a tough hydrogel, or a hydrogel-elastomer hybrid. In some configurations, the bio-adhesive layer 16 is configured to maintain high water content (>90 wt%) for the duration that the probe 10 is worn (e.g., up to 48 hours). The bio-adhesive layer 16 may also have a high interfacial toughness (e.g., >500 Jm’², or >850 Jm'²), and low acoustic attenuation (e.g., similar to commercial ultrasound gels). The bio-adhesive layer 16 can adhere the probe to the skin 22 of the user in order to continuously measure ultrasound shear wave elastography.

[0071] In some configurations, the probe 10 also includes a matching layer 24 arranged between to the bio-adhesive layer 16 and the piezoelectric layer 12. The matching layer 24 may include one or more layers of materials to compensate acoustic impedance mismatch between the probe and biological tissue (e.g., skin 22). For example, the matching layer 24 may include a layer of silver epoxy (e.g., 50 pm thick) and a layer of parylene (e.g., 45 pm thick).

[0072] In some configurations, the probe 10 may also include a backing layer 20. The backing layer may include alumina epoxy, a metal particle-polymer composites, or other materials.

[0073] Some configurations of the circuit layer, 18, bio-adhesive layer 16, matching layer 24, backing layer 20, and other components are further described below and in US20230277159, which is incorporated herein by reference in its entirety.

[0074] The probe 10 may have a small form factor to facilitate wearing on the skin. For example, in some configurations the probe 10 may have a length 30 in an azimuth dimension, which may be referred to as x, of <20 mm, 20-30 mm, 30-40 mm, 40-50 mm, 50-60 mm, 60-70 mm, 70-80 mm, 80-90 mm, or 90-100 mm. In some configurations, the length of the probe in each linear dimension is less than 30 mm, less than 40 mm, or less than 50 mm.

[0075] The wearable ultrasound elastography probe 10 can be used in systems that provide elastography measurements to evaluate stiffness of biological tissues under a subject’s skin. For example, stiffness of deep organs (e.g., 1-5 cm, 3-5 cm, 4-5 cm, up to 5 cm, up to 10 cm, or up to 20 cm under a subject’s skin) can be measured using the ARFI technique. In a refinement, the ARFI technique can be applied with a center frequency ranging from 2 to 8 MHz or 2 to 10 MHz. In another refinement, the ARFI technique can be applied with a center frequency of 7.5 MHz. Reducing the center frequency to 2-5 MHz may achieve a deeper penetration depth. The center frequency of the transducers can be chosen based on the desired resolution and penetration depth of the imaging and elastography measurements.

(0076] In some configurations, the wearable ultrasound elastography probe 10 may be used in a system that can additionally provide imaging of biological tissues. In this way, the system may advantageously provide functional maps or images of the subject’s organ of interest. For example, stiffness measurements can be displayed as a colormap overlaid onto an anatomical image of the liver. [0077] Referring now to FIG. IB, a flowchart is provided that sets forth steps of a process for manufacturing a wearable bio-adhesive ultrasound elastography probe. As a non-limiting example, the process 50 can be used to form all or part of the probe as shown and described with respect to FIG. 1 A. It should be understood that the steps of process 50 need not be performed in the order shown, as the layers can be coupled in any temporal order desired.

[0078] The process 50 includes forming slits in a layer of piezoelectric material, as shown in block 52. Such slits may be referred to as kerfs, grooves, or troughs. The slits define edges of discrete ultrasound transducer elements formed within the piezoelectric material. For example, N-l (e.g., 127, 255) slits can be formed in the piezoelectric material in order to provide N (e.g., 128, 256, respectively) ultrasound transducer elements.

[0079] As previously described, each slit can be described by a kerf size (e.g., part 34 in FIG. 1A). In some configurations the slits can be configured to have a kerf size constrained by Fs/4/c< 40 pm. As non-limiting examples, the kerf size may be configured as 20 pm, < 40 pm, < 30 pm, 10 pm- 30 pm, or 15 pm - 25 pm. The slits can also be defined based on a center-to-center separation, which similarly describes the center-to-center separation of the ultrasound transducer elements or pitch (e.g., part 32 in FIG. 1A). In some configurations, the center-to-center separation can be chosen based on the shear wavelength in the medium of interest (e.g., tissue or estimated as water). For example, the center-to-center separation can be configured as 200 pm, 150 pm-250 pm, 175 pm-225 pm, 195 pm- 205 pm, or 198 pm-202 pm.

[0080] The slits can be formed in the piezoelectric material using a scratch-dicing method. In some configurations the scratch-dicing can be performed using a dicing saw with a diamond blade, which may have a thickness less than the kerf size. In some configurations, block 52 may further include filling the slits with epoxy.

[00811 The layer of piezoelectric material is adhered to a matching layer, as in block 54. In block 56 a bio-adhesive layer can be implanted onto the matching layer or onto the ultrasound transducer. The bio-adhesive layer is configured to adhere to skin or a subject in order to facilitate the wearability of the ultrasound probe. Such layer may include a water-based tough hydrogel encapsulated by an elastomer membrane. The piezoelectric layer can also be bonded to a PCB or fPCB, as in block 58. In some configurations the PCB is bonded onto the piezoelectric layer on the opposite side from the matching and bio-adhesive layers.

[0082] Now referring to FIG. 1C, a process 60 of using a wearable bio-adhesive ultrasound probe is presented. In some configurations, process 60 can provide a method for measuring stiffness or imaging a tissue, such as a subcutaneous structure (e.g., liver). Process 60 may use an ultrasound system as described herein, for example, the system described with respect to FIG. 1A. In a refinement, the subject’s subcutaneous structure is within 5 cm of the subject’s skin. In a refinement, the subject’s subcutaneous structure is a liver or a portion thereof.

[0083] Process 60 includes placing a wearable ultrasound probe (e.g., shown in FIG. 1 A) onto the skin of a subject, as shown in block 62. Placing the probe on the skin may include adhering the probe onto the skin using the bio-adhesive layer. In some configurations the probe can remain on the skin for more than 24 or 48 hours. The process also includes controlling the ultrasound probe to produce an ARFI, as in block 64. Such ARFI can excite a shear wave that propagates through the biological tissue of the subject. In some configurations, the ARFI can be produced using a subset (e.g., center 32 elements) of the ultrasound transducer elements.

[0084] The shear wave causes particle displacement according to the stiffness of the tissue. The particle displacement can be measured by the ultrasound probe, as in block 66. In some configurations all of the ultrasound transducers can be used to measure the particle displacement. In other configurations, a subset of the ultrasound transducers can be used to measure the particle displacement. The particle displacement reflects tissue stiffness. Thus, the received ultrasound signal, which may be referred to as particle displacement data, elastography data, or shear wave elastography data. The elastography data can be processed by a processor to provide a measurement of tissue stiffness, as in block 68. (0085] The following examples illustrate the various embodiments of the present invention. Those skilled in the art will recognize many variations that are within the spirit of the present invention and scope of the claims.

[0086] Fabrication of wearable bio-adhesive ultrasound elastography (BAUS-E)

[0087] The development of the wearable ultrasound elastography device, BAUS-E, presented several design considerations to achieve optimal performance. To ensure good wearability, the device advantageously has a small size. Additionally, the device includes a sufficient number of channels for imaging, which includes 128 channels in a non-limiting example. Furthermore, the device can withstand high voltage driving to generate ARFI pushes, which pose fabrication challenges.

[0088] To meet these design goals, a 7.5 MHz wearable bio-adhesive ultrasound elastography device (BAUS-E) with a 128-element linear array was designed and fabricated as a non-limiting example. For the core component of the BAUS-E, a 1-3 composite lead zirconate titanate (PZT) piezoelectric material with both-side electrodes (e.g., Blatek Industries, Inc., Boalsburg, PA) was selected. The PZT material had a thickness of 140 pm, a pitch size of 100 pm, and a kerf size of 20 pm (Table 2, FIG. 12). These design specifications ensured that the width-to-thickness ratio was less than 0.7 which is a beneficial transducer design criterion. The width-to-thickness ratio was designed to be 0.57 (80 pm/140 pm). This ratio was chosen to generate a higher electromechanical coupling coefficient and enable pure thickness-mode resonance, resulting in better performance for ARFI-based elastography applications. A higher width-to-thickness ratio (> 0.7) could induce different resonance vibration modes that interfere with the fundamental resonance mode, which could reduce the bandwidth and signal-to-noise ratio (SNR). Therefore, maintaining the appropriate width-to-thickness ratio was important for achieving high device performance.

[0089] The transducer performance of the BAUS-E was simulated using a Krimboltz, Leedom, and Mattaei (KLM) equivalent circuit model in the PiezoelectricCAD software (Sonic Concepts Inc., Bothell, USA) to explore the optimization of the electrical impedance, phase spectrum, and pulse-echo performance of the BAUS-E prior to fabrication. Based on the simulation results, two quarterwavelength matching layers of the BAUS-E were designed and added on one side of the PZT composite to compensate acoustic impedance mismatch between the BAUS-E and biological tissues. These matching layers include silver epoxy and parylene. [0090] For the designed array, a pitch size of 1 Xwater (approximately 200 pm in water) was chosen to achieve higher imaging quality. The kerf size of the array element was set to 20 pm, and the composite kerf size was approximately 20 pm based on the criterion VJ f < 40 pm, where represents the shear velocity of the filled epoxy and/c is the center frequency of the array. Therefore, the top electrode was scratch-diced along the kerf direction of the PZT composite to form a 128- channel electrode separation in the array.

10091] The scratch-diced process was achieved by using a dicing saw (e.g., Tear 864-1, Thermocarbon, Casselberry, FL). Afterwards, the prepared composite component was bonded on a customized flexible printed circuit board (f-PCB) substrate. The alumina epoxy (e.g., Aluminum Oxide Powder, Buehler, IL, USA) was used to create a backing layer, which was adhered to the other side of the f-PCB to assemble a wearable ultrasound linear array. A bio-adhesive hydrogel was then implanted on the surface of the wearable ultrasound linear array to form a wearable array patch, BAUS-E. To further evaluate the BAUS-E, the electrical impedance and phase spectrum of each channel in the BAUS-E were measured using an Impedance Analyzer (e.g., Agilent 4294A, Santa Clara, CA). The resonant frequency (fr) and anti-resonant frequency (fa) can be obtained to calculate the electromechanical coupling coefficient (kt) using the following equation:

[0092] The pulse-echo performance, transducer sensitivity and bandwidth at -6 dB of the BAUS-E were tested using the Verasonics system (Vantage 256, Verasonics Inc. Kirkland, WA). By conducting these simulations and optimizing the transducer design parameters, the performance of the BAUS-E could be tailored to meet the specific goals of wearable ultrasound elastography.

[0093| Shear wave elastography (SWE) implementation

[0094] The BAUS-E was compatible with the Verasonics system for implementing SWEs in phantom and animal studies. The Verasonics system was driven by MATLAB scripts (MATLAB R2022a release, MathWorks, Natick, MA, USA) that generated the necessary structures to achieve the desired action. In the phantom study, both BAUS-E and a commercial linear array (L7-4, Philips, MA) were used to scan a commercial ultrasound phantom (Model 054GS, CIRS Inc. Norfolk, VA) for evaluating SWEs. The phantom provided a ground truth value of shear wave velocity as 2.8 m/s (Young's modulus E = 24 kPa provided by CIRS Inc.). The 32 elements in the middle of both transducers were used for the ARFI push, and all 128 elements were used for receiving particle displacements. The focal point of the ARFI for both devices was placed approximately 10 mm below the surface of the phantom, resulting in a /-number (/ ) of approximately 1.5 for the BAUS-E and 1.1 for the commercial array. In the phantom study using the commercial array, 1000 cycles of the ARFI push (push duration of 250 ps) were transmitted at 4.09 MHz, followed by plane wave compounding with three angles (-3°, 0°, +3°) performed immediately at 5 MHz for motion detection. To achieve a similar push duration in the phantom study using the BAUS-E, an ARFI push with 1875 cycles (push duration of 250 ps) was used. The ARFI push was transmitted at 7.5 MHz and the three angles of plane waves were immediately performed at the same frequency for motion detection. In the animal studies, the focus of the ARFI push of the BAUS-E was selected approximately 6 mm under the skin of the rats (near the middle of the liver). The ARFI push utilized the 20 elements in the middle of the BAUS-E with a push duration of 250 ps. Subsequently, all 128 channels were activated for motion detection. The f# of the BAUS-E in the animal study was approximately 1.5. In both the phantom and animal studies, the frame rate was set to 80 ps. Therefore, the effective frame rate after compounding was approximately 4.17 kHz.

[0095] For elasticity assessment, the in-phase/quadrature (IQ) data were saved for offline analysis. The autocorrelation was used to calculate the particle velocity from the IQ data, and the shear wave propagation was constructed via the spatiotemporal map. During the construction, a region of interest (ROI) was chosen from the axial displacement matrix. The ROI included the push row, five rows above the push, and the five rows below the push to average along the depth dimension over 0.5 mm. The maximum values of particle velocity at each time point were determined, and the shear wave group velocity C_g was obtained using a linear regression curve fitting method. To estimate the elastic properties of a biological tissue, simplifications of homogeneous, incompressible, and isotropic properties are commonly assumed although a real biological tissue represents a complex heterogeneous, anisotropic and non-linear behaviors. Therefore, the elasticity of the biological tissue can be treated as uniform distribution, and the tissue material response can be considered as nonorientation dependent. Under these assumptions, the C_g can be described as. c» ⁼ JI ⁽²⁾ where G is the shear modulus and p is the density of the medium, which was set as 1000 kg/m³ in the study. Due to the assumption of incompressible and isotropic properties, the relationship between shear modulus and Young’s modulus E can be expressed as

E = 3G (3)

Substituting equation (3) into (2), the relationship between C_g and E can be represented in the following way.

E = 3pc_g ² (4)

[0096] To further validate how accuracy of measuring C_g using BAUS-E in the phantom studies, Cg was transformed from the time domain to the frequency domain using a two-dimensional Fourier transform (2D-FT) to calculate wave dispersions. The 2D-FT was applied to the spatiotemporal map to obtain a frequency-wavenumber (f-k) domain distribution (k-space) of ultrasound shear waves. Shear wave phase velocity C_P were calculated by identifying the maximum peaks in the k-space using the equation Cp = Znflk, where /is frequency and k is wavenumber. A mean Cp over the frequencies from 400 to 1500 Hz on the dispersion curve was then computed. The aforementioned procedures were implemented using MATLAB R2022a software on a desktop computer with an Intel(R) Core(TM) i7-7700 CPU at 3.60 GHz processor, 16 GB memory and a 64- bit Windows 10 operating system.

[0097] Performance of wearable bio-adhesive ultrasound elastography (BAUS-E)

[0098] The proposed BAUS-E not only features a smaller active area for the comfortable wearability but also exhibits the capability to produce sufficient energy for ARFI push, allowing it to generate shear waves and provide the satisfied image quality for tracking particle displacements, as show in FIG. 2A. The BAUS-E was composed of an acoustic backing layer, a flexible printed circuit board (f-PCB), a 128-channel PZT composite, a matching layer, and a bio-adhesive hydrogel layer, as shown in FIG. 2B.

[0099] The bio-adhesive couplant can be composed of a 95 wt% water-based tough hydrogel encapsulated by an elastomer membrane. A thin bio-adhesive layer was further coated to maintain a robust adhesion over 48 hours. FIG. 2C demonstrates an example of a bio-adhesive ultrasound probe adhered to skin of a subject, demonstrating its ability to withstand a high pulling force and maintain robust adhesion. The scale bar indicates 10 mm. [0100] An electrode separation by a dicing saw was imaged using optical microscopy to show a 128-channel BAUS-E within only 24 mm aperture size, as shown in FIG 2B where the scale bar indicates 10 mm. In FIG. 12, Table 2 provides parameters for a non-limiting example BAUS-E. The electrode separation of each element was finely diced with the 20 pm diamond blade by a dicing saw to achieve 128 individual channels, as demonstrated in FIG. 2B, for example. The center frequency of the BAUS-E was designed to be 7.5 MHz to strike a balance between the expected spatial resolution and the required energy for ARFI pushes. The 1 Xwater pitch size (approximately 200 //m) and 20 z/m kerf size were selected based on the optimized thickness-vibration mode theory, which provided the active area of the BAUS-E to have approximately 4 mm in elevational direction and merely 24 mm in azimuth direction. This size provided a suitable size to achieve both a sufficient field-of-view (FOV) for shear wave propagations and comfortable wearability.

|O101| To characterize the performance of the BAUS-E, the pulse-echo response, peak-peak sensitivity, -6dB fractional bandwidth, electrical impedance, lateral resolution and axial resolution were measured. For the pulse-echo response, the fiftieth channel was randomly selected, and its center frequency was measured at 7.2 MHz (FIG. 3A). Comparing with the PiezoelectricCAD simulation result of the center frequency of 7.5 MHz, there is a good agreement between the measured and simulated center frequencies of the BAUS-E (FIG. 3B). The mean center frequency of the BAUS-E across all 128 channels was determined to be 7.5 ± 0.26 MHz (FIG. 3C). The electrical impedance and phase spectrum were also evaluated and compared with the simulation results. The impedance of the fiftieth channel was measured to be approximately 165 Ohms (FIG. 4A), which aligns well with the simulation result of 164 Ohms (FIG. 4B). The mean impedance of the BAUS-E across all 128 channels was found to be 164.72 ± 2.76 Ohms (FIG. 4C), demonstrating close agreement with the simulation results.

[0102] To further assess the performance of the BAUS-E, we evaluated the peak-peak sensitivity, -6 dB fractional bandwidth, lateral resolution and axial resolution. The mean peak-peak sensitivity was determined to be -4.14 ± -2.36 dB for the BAUS-E and -3.17 ± -1.51 dB for the commercial transducer (FIG. 5A). The peak-peak sensitivity reflects the ability of an ultrasonic device to detect reflectors at a specific depth in a test material. A value closer to 0 dB indicates higher sensitivity of the ultrasonic device. Regarding the -6 dB fractional bandwidth, the mean fractional bandwidth was 37.44 ± 4.85% for the BAUS-E and 44.34 ± 3.05% for the commercial transducer (FIG. 5B). The BAUS-E exhibited a comparable performance to the commercial transducer in terms of peak-peak sensitivity and -6 dB fractional bandwidth. For the resolution of the BAUS-E in B-mode imaging, the lateral resolution and the axial resolution, determined by the -6 dB full-width halfmaximum (FWHM) beam profile, were approximately 0.13 mm (target @ 13 mm) and 0.36 mm using tungsten wire phantom (FIG. 6A), respectively (FIG. 6B). Theoretical calculations predicted lateral and axial resolutions of 0.14 mm and 0.39 mm, respectively, based on the BAUS-E's center frequency of 7.5 MHz and aperture size of 24 mm (FIG. 6C). Our results demonstrate that the BAUS-E device exhibits performance close to the simulation results of the transducer design and is comparable to the commercial ultrasound transducer.

[0103] Numerical simulations

10104] To validate the proposed BAUS-E technique, numerical simulations were conducted using a staggered grid finite difference method (SGFD). The SGFD scheme was implemented and used to generate particle velocity shear wave motion data. The SGFD scheme utilized the velocitystress first-order hyperbolic system of equations. To minimize undesired reflections from model boundaries, a complex-frequency shifted perfectly matched layer based on recursive integration was implemented. The Field II software package was utilized to simulate the ARFI push beam. For the validation in the phantom studies, a two-dimensional model with the elastic material property of the numerical phantom was considered with 23 mm width x 20 mm thickness for the BAUS-E and 40 mm width x 20 mm thickness for the L7-4 transducer. The elastic domains were uniformly sampled at a spatial resolution of 0.1 mm. The thickness of the bio-adhesive couplant was set to 1.2 mm based on measurements. Based on experimental results obtained using BAUS-E, the C_g in the numerical phantom was set to 2.8 m/s. The Young’s modulus E of the numerical phantom can be approximately evaluated by the formula 3pC_g ² and was set to 24 kPa as the same value in the CIRS phantom, where Cg is the shear wave velocity, and the density of the materials p was set to 1000 kg/m³ in the simulation. To consider the boundary condition with the bio-adhesive couplant, E was set to 100 kPa (corresponding to an approximate C_g of 5.77 m/s) as previously reported. The simulation time step was set to 0.03 ps to ensure the stability of the numerical model. The ARFI excitation with a pulse duration of 250 /zs was applied in the axial direction for both the commercial ultrasound transducer and the BAUS-E in the phantom study (FIG. 7). (0105] In particular, FIG. 7 provides a set of correlated graphs. The first two rows show the results using the commercial transducer, which compare with the last two rows for the results using the BAUS-E device. Graph 700 shows the simulated ARFI push using the commercial transducer without the bio-adhesive couplant and graph 702 shows it with the bio-adhesive couplant. Graph 704 shows the simulated ARFI push using BAUS-E device without the bio-adhesive couplant and graph 706 shows it with the bio-adhesive couplant. The bio-adhesive couplant was indicated as the red arrow. Graphs 708, 710, 712, and 714 illustrate the simulated shear wave velocity at 10 mm under the surface of the numerical phantom. The simulated shear wave velocity measured by the BAUS-E shows a good agreement with the results using the commercial transducer and the bio-adhesive couplant would not affect the ARFI push and shear wave measurements, which has been demonstrated in the numerical simulations. Graphs 716, 718, 720, and 722 show the k-space based on the information of graphs 708- 714 (rectangle formed by dashed line) using 2D-FT. Graphs 724, 726, 728, and 730 exhibit the dispersion curves of graphs 708-714 to evaluate the shear wave phase velocities and compared them with their theoretical value (2.80 m/s). The shear wave phase velocities using the commercial transducer and BAUS-E w/wo bio-adhesive couplant show a good agreement with the theoretical value. BC: bio-adhesive couplant; Comm: commercial transducer.

[0106] The focus of the ARFI push was set at 10 mm below the surface of the numerical phantom. The simulations were performed using parallel computation technology with modem graphics processing units (GPUs) and compute unified device architecture (CUD A). The calculations were conducted in MATLAB R2022a release on a desktop computer equipped with an Intel(R) Core(TM) i7-6700 CPU at 3.40 GHz processor, 48 GB memory and 64-bit Windows 10 Pro operating system, along with an NVIDIA GeForce RTX 2080 Ti graphics card. The computation time for the numerical simulations was approximately 25 minutes.

[0107] Fabrication of polyacrylamide hydrogel-elastomer hybrid bio-adhesive couplant [0108] The BAUS-E couplant contains a soft and tough hydrogel encapsulated by a bioadhesive elastomer membrane to form a hydrogel-elastomer hybrid. The PAAm-Chitosan tough hydrogel was synthesized by UV photo-polymerization of 3% (w/w) high molecular weight chitosan, 12% (w/w) acrylamide, 0.1% (w/w) N, N'-methylenebisacrylamide, and 0.2% (w/w) a-ketoglutaric acid in a 1% (v/v) acetic acid solution. The obtained tough hydrogels were immersed in IM CaC12 solution for 24 hours to reach an equilibrium state. The bio-adhesive membrane was designed as a sandwich structure with two layers of bio-adhesive polyurethane and one layer of hydrophobic polyurethane. The bio-adhesive polyurethane was synthesized by grafting polyacrylic acid chains and polyethylhexyl acrylate chains onto hydrophilic polyurethane. We used 30% (w/w) acrylic acid and 5% (w/w) 2-ethylhexyl acrylate to modify the hydrophilic polyurethanes. After the reaction, the mixture solution was dialyzed (Cutoff Mn 3000 Da) against ethanol for 3 days and against water for 3 more days to obtain the bio-adhesive polyurethane. A thin layer of bio-adhesive polyurethane with 5% (w/w) EDC and 5% (w/w) NHS was spin-coated at 1500 rpm on a clean glass, and then hydrophobic polyurethane was spin-coated at 1500 rpm on the bio-adhesive polyurethane. Before the hydrophobic polyurethane layer was fully dried, another thin layer of bio-adhesive polyurethane with 5% (w/w) EDC and 3% (w/w) NHS was spin-coated at 1500 rpm on the hydrophobic polyurethane layer. The film was then placed in airflow for 4 hours for drying to obtain the bio-adhesive membrane. To fabricate the hydrogel-elastomer hybrid bio-adhesive couplant, the hydrogel was tailored into the desired shape and size, and the adhesive membrane was adhered onto the hydrogel with a gentle press to avoid any bubbles. The BAUS-E transducer was firmly adhered to the couplant using a layer of 50% (w/w) hydrophilic polyurethanes in a 70% (v/v) ethanol and 30% (v/v) water solution. The fabrication of the bio-adhesive hydrogel-elastomer hybrid couplant for the BAUS-E followed the same procedure as the BAUS couplant as described in US20230277159A1.

[0109) Systematic phantom evaluations of the elasticity and numerical simulations of BAUS-E

101101 To evaluate the performance of the BAUS-E in measuring mechanical properties using SWE, a systematic phantom study was conducted (FIG. 8). In particular, graphs 800, 802, 804, and 806 show shear wave measurements in the time domain using different experimental conditions, where: graph 800 was generated using a commercial ultrasound transducer coupled with clinical ultrasound gel, graph 802 was generated using a commercial ultrasound transducer coupled with bioadhesive couplant, graph 804 was generated using a proposed wearable device with clinical ultrasound gel, and graph 806 was generated using the proposed wearable device with bio-adhesive couplant, the BAUS-E. Graphs 808, 810, 812, and 814 show numerical simulation results of shear waves corresponding to the four different experimental conditions, demonstrating good agreement with the experimental results. Graphs 816, 818, 820, and 822 provide data from a 2D-FT analysis of the shear wave measurements for graphs 800-806. Graph 824 shows mean shear wave velocities (n = 10) for the four different measurements presented with their standard deviations. Graph 826 shows dispersion curves for that of graph 800 and graph 802 in the frequency range of 400 to 1500 Hz compared with the numerical simulation results. Graph 828 shows dispersion curves for that of graph 804 and graph 806 in the frequency range of 400 to 1500 Hz compared with the numerical simulation results. Specifically, a commercial ultrasound phantom was used in this study to provide a ground truth value of shear wave velocity (Cg, group velocity), which was determined to be of 2.80 m/s provided by CIRS Inc. Based on Eq. (4), the Young's modulus was calculated as 23.52 kPa, assuming a density of 1000 kg/m³ (provided by CIRS Inc.). This systematic experiment was divided into four different experimental conditions: SWE measured by the commercial ultrasound transducer with the conventional ultrasound coupling gel (control), SWE measured by the commercial ultrasound transducer interfaced with the bio-adhesive couplant, SWE measured by the wearable device with the conventional ultrasound coupling gel, and SWE measured by the BAUS-E. To ensure consistency, the focus of the ARFI for both devices were placed at 10 mm below the phantom surface. The C_g measured by the commercial ultrasound transducer with the conventional ultrasound coupling gel and with bioadhesive couplant was 2.80 m/s (represented by the dashed line, FIG. 8, graph 800) and 2.82 m/s (FIG. 8, graph, 802), respectively. The C_g was 2.81 m/s (FIG. 8, graph 804) measured by the wearable ultrasound device with the ultrasound coupling gel and 2.78 m/s by the BAUS-E (FIG. 8, graph 806). The mean C_g (n =10) measured by the commercial transducer with the conventional ultrasound coupling gel, the commercial transducer with the bio-adhesive couplant, the wearable device with the conventional ultrasound coupling gel and the BAUS-E was 2.81 ± 0.01 m/s (control), 2.82 ± 0.01 m/s, 2.79 ± 0.01 m/s and 2.78 ± 0.04 m/s, respectively (FIG. 8, graph 724). Based on Eq. (4), the Young's modulus ranged between 23.18 kPa and 23.85 kPa, demonstrating good agreement with the ground truth value of 23.52 kPa. The systematic phantom experiments demonstrated that the BAUS-E exhibited good performance for SWE. Furthermore, the phantom studies showed that the bio-adhesive couplant did not interfere the results of SWE due to its good acoustic impedance matching.

[OiH I Numerical simulations were further utilized to validate the group velocity of shear waves in the four experimental cases as the secondary validation (FIG. 7, graphs 700-706). The simulated ARFI focus was placed at a depth of 10 mm under the surface of the numerical phantom, mimicking the experimental setup. The ground truth value of the elastic property in the numerical phantom was set to 23.52 kPa, which matched the Young's modulus of the CIRS commercial phantom. The elastic property of bio-adhesive couplant was set to 100 kPa based on our previous research. The simulated C_g was 2.802 m/s for the commercial ultrasound transducer with the conventional ultrasound coupling gel (control) (indicated by the dashed line, FIG. 8, graph 808; FIG. 7, graph 708), and 2.803 m/s for the transducer with the bio-adhesive couplant (FIG. 8, graph 810; FIG. 7, graph 710). The simulated C_g was 2.798 m/s measured by the wearable device with the ultrasound coupling gel (FIG. 8, graph 804; FIG. 7, graph 712) and 2.803 m/s for the BAUS-E (FIG. 8, graph 806; FIG. 7, graph 714). According to the Eq. (4), the Young’s modulus ranged between 23.35 kPa and 23.52 kPa, showing high consistency with the ground truth value of 23.52 kPa set in the numerical phantom. The results of the numerical simulations exhibited strong agreement with the experimental results across the four different experimental conditions for shear wave velocity measurements.

|O112] Furthermore, we evaluated the dispersion curves of shear waves in the phantom study to analyze the frequency components of the measured shear waves in the time domain. The shear waves in the spatiotemporal map (indicated by the dashed rectangle, FIG. 8, graphs 800-814) were transformed into the frequency domain to calculate the k-space, which is a function of frequency (Hz) and wavenumber (1/m), allowing us to evaluate the shear wave velocity (C_P, phase velocity). The k- spaces from the experimental results (FIG. 8, graphs 816-822) and the numerical simulation results (FIG. 7, graphs 716-722) were obtained using 2D-FT. Based on the k-space information, the mean wave dispersion curves were calculated for the experimental and the numerical simulation results (FIG. 8, graphs 826 and 828, FIG. 7, graphs 724-730). The mean C_P with the frequency range of 400 to 1500 Hz was determined to be 2.84 ± 0.01 m/s (corresponding to 24.19 kPa) (FIG. 8, graph 826) for the commercial ultrasound transducer with the conventional ultrasound coupling gel, and 2.82 ± 0.01 m/s (corresponding to 23.85 kPa) (FIG. 8, graph 826) for the transducer with the bio-adhesive couplant. In the numerical simulation with the commercial ultrasound transducer, the mean C_P with the conventional ultrasound coupling gel was 2.81 ± 0.01 m/s (indicated by the dot, FIG. 8, graph 826; FIG. 7, 724) and with the bio-adhesive couplant, it was 2.82 ± 0.01 m/s (corresponding to 23.68 kPa) (FIG. 7, graph 726). On the other hand, for the wearable device with the conventional ultrasound coupling gel, the mean C_P was 2.76 ± 0.02 m/s (corresponding to 22.85 kPa) (FIG. 8, graph 828), and for the BAUS-E, it was 2.74 ± 0.02 m/s (corresponding to 22.52 kPa) (FIG. 8, graph 828). Comparing with the numerical simulations for the case with the wearable device, the mean C_P with the conventional ultrasound coupling gel was 2.80 ± 0.01 m/s (corresponding to 23.52 kPa) (dot, FIG. 8, graph 828; FIG. 7, 728) and 2.81 ± 0.01 m/s (corresponding to 23.68 kPa) for the BAUS-E (FIG. 7, graph 730). According to these three validation methods, the BAUS-E demonstrated satisfactory performance in ARFI-based wearable ultrasound shear wave elastography.

[0113] The ARFI pushes in commercial ultrasound systems are typically performed at a distance of 4-4.5 cm from the transducer to achieve optimal liver stiffness measurements in clinical settings. However, for most wearable ultrasound devices, assessing the elastic properties of deep biological tissues remains challenging. This is because achieving a smaller £#, which is necessary for a smaller aperture size to ensure comfortable wearability, can result in a weaker push force at deeper tissue regions. While mechanical compression-based methods have been reported to evaluate tissue stiffness up to 4 cm beneath the skin, larger loading forces may be required to avoid wave distortions when using stretchable ultrasound arrays. Consequently, assessing the elasticity of internal organs such as the liver could be difficult. Furthermore, mechanical compression methods often require additional excitation tools such as fingers or actuators, which makes continuous measurements challenging especially for the diseases that rapidly change. In the study, we performed shear wave measurements at various depths using the BAUS-E (FIG. 9).

|0114| In particular, FIG. 9 shows the shear wave measurement in various depths in the phantom test. In graphs 900, 902, 904 and 906, the simulated ARFI push at 20 mm, 30 mm, 40 mm and 50 mm depth under the surface of the phantom are shown, respectively. Graphs 908, 910, 912, and 914 exhibit the shear wave propagation (arrow) at 20 mm, 30 mm, 40 mm and 50 mm depth under the surface of the phantom, respectively. Graphs 916, 918, 920, and 922 illustrate the spatiotemporal map of the shear wave velocity at 20 mm, 30 mm, 40 mm and 50 mm depth under the surface of the phantom, respectively.

[0115] The phantom test involved ARFI pushes at depths of 20 mm, 30 mm, 40 mm, and 50 mm beneath the surface of the phantom (FIG. 9, graph 900-906) and consistent results of C_g, approximately 2.8 m/s, were obtained (FIG. 9, graphs 916-922). Despite having a smaller aperture size for comfortable wearability, the BAUS-E demonstrated the capability to measure elasticity up to 50 mm (FIG. 9, graph 916). This highlights the superior performance of the BAUS-E to reach similar conditions as ARFI pushes performed in clinical settings, thus demonstrating its potential for translational applications. (0116] Liver stiffness using BAUS-E reflects the severity of acute liver failure (ALF) in rats

[0117| The animal model with pharmacological induction of ALF using D-Galactosamine was employed. The BAUS-E was utilized to continuously monitor the stiffness changes of rat livers with ALF over 48-hour period with measurements taken at 6-hour intervals, as show in FIG. 10A. Thus, the liver stiffness changes in rats with ALF were measured at 9 different time points. The experiments were repeated five times (n = 5), and each measurement at every time point consisted of three shear wave measurements, resulting in a total of 15 measurements at each time point. The C_g of rat livers at four different time points representing liver injuries (0, 12, 30 and 48 hours) and measured using the BAUS-E were 1.23 m/s, 1.53 m/s, 1.80 m/s and 2.42 m/s, respectively (FIG. 10B). The mean C_g of normal livers in rats (control) was 1.17 ± 0.08 m/s. At 6 hours after the injection of D-Galactosamine, the mean C_g increased to 1.38 ± 0.05 m/s. Between the 12th and 30th hours, the mean C_g in rat livers with ALF gradually increased from 1.56 ± 0.05 m/s, 1.61 ± 0.07 m/s, 1.71 ± 0.09 m/s to 1.82 ± 0.09 m/s, respectively. A significant increase in liver stiffness was observed between the 30th and 36th hours with the mean C_g in rat livers with ALF increasing to 2.48 ± 0.12 m/s. After the 36th hour, the mean C_g reached a plateau around 2.5 m/s (FIG. 10C). There was a statistically significant difference in liver stiffness between normal livers and livers with ALF at each time point measurements (p < 0.001) (FIG. 10C and D). The mean Young's modulus of normal livers in rats was 4.16 ± 0.55 kPa, which aligns well with the shear modulus measurements of 1.50 ± 0.10 kPa reported in well- established literature. It's worth noting that the Young's modulus can generally be assumed as three times the shear modulus according to Eq. (3). The mean Young's modulus of livers after the injection of D-Galactosamine gradually increased from 5.70 ± 0.44 kPa, 7.35 ± 0.48 kPa, 7.82 ± 0.74 kPa, 8.77 ± 0.92 kPa to 10.00 ± 0.97 kPa in the SWE measurements between the 6th and 30th hours. The Young's modulus rapidly increased to 18.47 ± 1.88 kPa at the 36th hour, and liver stiffness remained stable until the 48th hour (18.08 ± 2.62 kPa at 42 hours and 18.17 ± 2.44 kPa at 48 hours) (FIG. 10D). These results demonstrated that using the BAUS-E can effectively measure the liver elasticity in rats with ALF as a valuable biomarker and can reflect the progression of ALF through rapid changes observed in continuous moduli measurements.

[0118] Example system (0119] Referring to FIG. 13, one non-limiting example of an electronic or sensor device in accordance with the present disclosure is provided. In this example, the electronic device is illustrated as an ultrasound system 100 that can implement the methods described in the present disclosure is shown. The ultrasound system 100 includes a transducer array 102 that includes a plurality of separately driven transducer elements 104. The transducer array 102 can include any suitable ultrasound transducer array, including linear arrays, curved arrays, annular array, phased arrays, and so on. Similarly, the transducer array 102 can include a ID transducer, a 1.5D transducer, a 1.75D transducer, a 2D transducer, a 3D transducer, and so on.

[0120] When energized by a transmitter 106, a given transducer element 104 produces a burst of ultrasonic wave. The ultrasonic wave reflected back to the transducer array 102 (e.g., an echo) from the object or subject under study is converted to an electrical signal (e.g., an echo signal) by each transducer element 104 and can be applied separately to a receiver 108 through a set of switches 110. The transmitter 106, receiver 108, and switches 110 are operated under the control of a controller 112, which may include one or more processors. As one example, the controller 112 can include a computer system. In a non-limiting example, the controller 112 may wirelessly connect to the transmitter 106, receiver 108, and switches 110. Alternatively, the controller 112 may connect to the transmitter 106, receiver 108, and switches 110 via wired connection. In a non-limiting example, the controller 112 may be portable.

[0121] The transmitter 106 can be programmed to transmit unfocused or focused ultrasound waves. In some configurations, the transmitter 106 can also be programmed to transmit diverged waves, spherical waves, cylindrical waves, plane waves, or combinations thereof. Furthermore, the transmitter 106 can be programmed to transmit spatially or temporally encoded pulses. In some configurations, the transmitter 106 can be programmed to transmit ARFI to produce shear waves.

[0122] In some configurations, beam steering may be used. Beam steering may be accomplished by energizing each transducer element 104, or a subset of transducer elements, in a specified sequence to steer an ultrasound beam along path R to target point P at an angle 0 from the axis of the transducer array 102.

[0123] The receiver 108 can be programmed to implement a suitable detection sequence for the elastography or imaging task at hand. In some embodiments, the detection sequence can include one or more of line-by-line scanning, compounding plane wave imaging, synthetic aperture imaging, and compounding diverging beam imaging.

[0124] In some configurations, the transmitter 106 and the receiver 108 can be programmed to implement a high frame rate. For instance, a frame rate associated with an acquisition pulse repetition frequency (“PRF”) of at least 100 Hz can be implemented. In some configurations, the ultrasound system 100 can sample and store at least one hundred ensembles of echo signals in the temporal direction.

[0125] The controller 112 can be programmed to design an imaging or elastography sequence using the techniques described in the present disclosure, or as otherwise known in the art. In some embodiments, the controller 112 receives user inputs defining various factors used in the design of the imaging or elastography sequence.

[0126] A scan can be performed by setting the switches 110 to their transmit position, thereby directing the transmitter 106 to be turned on momentarily to energize transducer elements 104 during a single transmission event according to the designed imaging or elastography sequence. The switches 110 can then be set to their receive position and the subsequent echo signals produced by the transducer elements 104 in response to one or more detected echoes are measured and applied to the receiver 108. The separate echo signals from the transducer elements 104 can be combined in the receiver 108 to produce a single echo signal.

[0127] The echo signals are communicated to a processing unit 114, which may be implemented by a hardware processor and memory, to process echo signals, images generated from echo signals, or elastography data generated from echo signals. As an example, the processing unit 114 can process images or elastography data from a wearable system using the methods described in the present disclosure. Images and elastography data or elastography maps produced from the echo signals by the processing unit 114 can be displayed on a display system 116.

[0128] In some configurations, any suitable computer-readable media can be used for storing instructions for performing the functions and/or processes described herein. For example, in some configurations, computer-readable media can be transitory or non -transitory. For example, non- transitory computer-readable media can include media such as magnetic media (e.g., hard disks, floppy disks), optical media (e.g., compact discs, digital video discs, Blu-ray discs), semiconductor media (e.g., RAM, flash memory, EPROM, EEPROM), any suitable media that is not fleeting or devoid of any semblance of permanence during transmission, and/or any suitable tangible media. As another example, transitory computer-readable media can include signals on networks, in wires, conductors, optical fibers, circuits, or any suitable media that is fleeting and devoid of any semblance of permanence during transmission, and/or any suitable intangible media.

(0129] Example wearable materials and configurations

[0130] The following description provides systems and methods for securing electronic devices directly on-person in a way that facilitates extended wearing, while properly positioning and adjusting the electronic device(s) for effective, extended, data gathering and/or ultrasound wave production.

[0131] In particular, an elastomer membrane may be utilized and coupled to the electronic device. An electronically-communicable or transmissible material may be arranged within the elastomer membrane. A bio-adhesive layer may be coupled to the elastomer membrane on a side opposite to the electronic device and configured to couple the elastomer membrane to the subject.

[0132] In one non-limiting example the electronic device may be an ultrasound transducer or array of ultrasound transducers. Furthermore, the electronic device may include a processor and/or communications components. In this way, the systems and methods described below can be used in accordance with the present disclosure to effectively adhere an ultrasound system to a biological tissue, such as the skin. The ultrasound system can be used to produce ARFI that excite shear waves in a biological tissue to provide ultrasound elastography that characterizes tissue stiffness. The ultrasound system can also be used to generate anatomical images.

|0133] As a particular but non-limiting example, systems and methods are provided for high- resolution wearable ultrasound imaging and/or elastography. In some configurations, a bio-adhesive ultrasound (BAUS) device may be used, which includes a thin, rigid, and high-resolution ultrasound probe robustly adhered on a subject, such as the skin of a subject, via a couplant layer made of a soft, tough, anti-dehydrating, and bio-adhesive hydrogel-elastomer hybrid material. A thin rigid probe provides for high transducer density, stable transducer positions under dynamic body motions, high reliability of the probe in long-term applications, and the like. A probe integrated on couplant may include a thin rigid ultrasound probe, handheld ultrasound probe, stretchable ultrasound probe, and the like. (0134] Referring to FIG. 14, a non-limiting example on-person, wearable system 200 is shown. Again, in this non-limiting example, the electronic device or sensor is an ultrasound probe 202. As previously described (e.g., in reference to FIG. 1 A), the ultrasound probe 202 may include an array of transducer elements 204. In some configurations, the ultrasound probe 202 may be a thin, rigid or flexible ultrasound probe, such as by including a substrate 205, which may be a circuit board, f-PCB, or other such structure. The wearable system 200 may include a couplant 206. In some non-limiting examples, the couplant 206 may be a hydrogel-elastomer hybrid with bio-adhesive layers. While conventional systems may require the skin to deform to the rigidity of the ultrasound array, or for the ultrasound array to be bent to conform to the skin, the couplant 206 may be used to conform to the skin surface 208 without requiring either the skin surface 208 to deform or to require the ultrasound probe 202 to bend or deform.

|O135| In some configurations of a hydrogel-elastomer hybrid, the hydrogel may include hydrophilic polymers or copolymers that have good acoustic transmission. In non-limiting examples, the hydrogel may include a material such as polyacrylic acid, polyacrylamide, polyvinyl alcohol, polyhydroxy ethyl methacrylate, polyethylene glycol, polyurethane, casein, albumin, gelatin, chitosan, hyaluronic acid, alginate, oxidized alginate, cellulose, oxidized cellulose, K-Carrageenan, sulfonated polysaccharides and the like. In some configurations, the hydrogel may include chemical or physical crosslinkers (e.g., gelatin methacrylate, hyaluronic acid methacrylate, oxidized methacrylic alginate, poly caprolactone diacrylate, N,N’-bis(acryloyl)cystamine, N,N’-methylenebis(acrylamide), polyvinyl alcohol, acrylic cyclodextrin(CD), and adamantlyacrylate, t-butylacrylate, polyethylene glycol diacrylate, polyethylene glycol dimethacrylate), in a single, double, or multiple networks.

[0136] The elastomer may include one, or more elastic polymers or rubbers that have high stretchability. Non-limiting example elastomer material include natural rubbers, styrene-butadiene block copolymers, polyisoprene, polybutadiene, ethyl ene-propylene rubber, ethylene propylene diene rubber, silicone elastomers, fluoroelastomers, polyurethane elastomers, nitrile rubbers, and the like. [0137] The bio-adhesive layer may include of one or more adhesive polymers that can adhere to the skin without corrosions. Non-limiting example bio-adhesive layer materials include animal glue, casein glue, starch, dextrin, agar, algin, gum arabic, epoxy resins, nitrocellulose, polyvinyl acetate, vinyl acetate-ethylene copolymer, polyethylene, polypropylene, polyamides, polyesters, acrylics, cyanoacrylics, natural rubber, butyl rubber, butadiene rubber, styrene-butadiene rubber, nitrile rubber, silicone, neoprene, and the like.

[0138] In a non-limiting example, the hydrogel-elastomer couplant includes polyacrylamide (PAAm) and chitosan interpenetrating hydrogel, polypropylene elastomer, and poly(2-ethylhexyl acrylate-co-acrylic acid) bio-adhesive.

[0139] In another non-limiting example, the hydrogel-elastomer couplant includes polyacrylamide (PAAm) and K-Carrageenan interpenetrating hydrogel, and polyurethane bio-adhesive elastomer.

[0140] Referring to FIG. 15, a cross-section of a non-limiting example wearable ultrasound system is shown. A thin, rigid ultrasound probe 302 may be coupled to a couplant 303 via an upper bio-adhesive layer 304. The couplant 303 may include an elastomer membrane 306 and 310 encapsulating a hydrogel 308. A lower bio-adhesive layer 312 may be used to provide fixed, but removable coupling to a skin surface 314. Ultrasound transmission 316 may propagate from the probe

302 to the skin 314 on transmit, and ultrasound may be detected from the skin 314 on detection.

[0141] The couplant 303 may also include the hydrogel 308, which, like the elastomer membrane 306, 310 is designed for transmission of electronic signals or sensing information from the skin 314 and tissue or organs below the skin 314. In one non-limiting example, the hydrogel 308 may contain 95 wt% water and be encapsulated by the elastomer membrane 308 and 310 and further coated by a thin bio-adhesive layer with upper and lower bio-adhesive layers 304 and 312. The couplant 303 may maintain robust adhesion between the ultrasound probe 302 or probe and the skin 314 over the long term and may insulate the ultrasound probe 302 from skin deformation during dynamic body motions. For example, the system may be designed to be worn for multiple days, for example, 48 hours or more.

[0142] The couplant 303 may provide a set of characteristics and functions inaccessible to common ultrasound couplants. The elastomer membrane 306, 310 may be used to prevent dehydration of the encapsulated hydrogel and ensures robust and comfortable adhesion of the probe on the skin over the long term, such as >4 hours, > 12 hours, or >48 hours. In some configurations, the couplant

303 may be soft, and tough couplant that maintains elasticity and is stable under high forces. The couplant 303 may also shield the probe from skin deformation. [0143] The acoustic transmissivity of the couplant 303 may maintain high transmissivity over the long term. In a non-limiting example, the acoustic transmissivity is >97% transmissivity relative to degassed water, and long term. In some configurations, the couplant 303 can readily adhere thin- profile probes of various commercial ultrasound devices to the skin over the long term. In some configurations, the couplant 303 may be fully constituted of common chemicals, which may be amenable to mass production as a low-cost medical supply and facilitate the broad applications of devices. In a non-limiting example, the total thickness of the elastomer membrane and the bio-adhesive layer may be below of the ultrasound wavelength.

[0144] The wearable ultrasound system may also include an optional heater 320 to provide thermal energy 322 to the hydrogel 308. The heater 320 may be removably coupled to the wearable ultrasound system. The thermal energy 322 can heat the hydrogel 308 of the adjustable BAUS couplant 303 from a normal skin temperature (e.g., 35°C) to an elevated yet skin-tolerable temperature (e.g., 50°C) for in situ adjustment of the adjustable BAUS probe’s position on the skin.

[0145] In some configurations, the ultrasound probe includes piezoelectric transducers with a center frequency ranging from 2 MHz to 10 MHz. The center frequency of the transducers determines the resolution and the penetration depth of the imaging and elastography measurement, which can achieve a resolution of 200 pm for a penetration depth less than 6 cm and a resolution of 600 pm for a penetration depth up to 18 cm. Each transducer may be controlled by the PCB 18.

[0146] The circuit layer may include ultrasound circuits. The circuits can be may be fabricated with any appropriate technique, such as three-dimensional (3D) printing, laser etching, photolithography, and the like. In some configurations, circuit-line resolutions of 100 pm, 10 pm, and 1 pm, may be achieved for each method, respectively. In some configurations, the optional bottom circuit may be covered by an acoustic matching layer to enhance the acoustic transmissivity to the skin. The top circuit may be covered by an acoustic backing layer to quench any resonance effect.

[0147] In some configurations, the ultrasound probe may be sealed by a layer of epoxy for high stability and reliability in long-term applications. To connect the ultrasound probe to an image or elastography processing system, a wireless connection may be used, or a “plug-and-play” input/output (VO) may be used for immediate download of data of the probe. In a non-limiting example, an I/O may include a flexible flat cable for forming a connection. (0148] An ultrasound probe may be sized for use on a subject. In a non-limiting example, the probe may have a thickness of <1 cm (e.g., 3 mm, 4 mm), and a length and width ranging from 2 cm to 4 cm. The ultrasound probe may have a much smaller size and lighter weight (e.g., l-10g, or 10-40 g) than conventional ultrasound imaging probes.

(0149] The bottom surface of the ultrasound probe (i.e., the matching-layer side) may be robustly adhered to the skin via the bio-adhesive couplant. The couplant may include a soft, yet tough hydrogel composed of chitosan-polyacrylamide (PAAm) interpenetrating polymer networks (such as 10 wt%) and water (such as 90 wt%). The hydrogel may be encapsulated by a thin elastomer membrane, such as with a thickness, < 40 pm, of polyurethane to prevent dehydration of the hydrogel and to make the skin comfortably contact with a dry couplant surface, instead of a wet hydrogel, over a long term (>48 hours). The polyurethane may be grafted with poly(acrylic acid) coupled with N- hydroxysuccinimide ester (NHS ester) to form robust bonding between the elastomer membrane and the hydrogel. The hydrogel-elastomer hybrid may be further coated by a thin bio-adhesive layer (such as with a thickness, <10 pm) synthesized by copolymerizing poly(ethylene glycol) diacrylate, 2- ethylhexyl acrylate, and acrylic acid.

|0150| The carboxylic acid, ethyl, and hexyl groups in the bio-adhesive layer may form physical bonds, such as hydrogen bonds and electrostatic interactions with the skin and the probe surface, providing instant, stable, and noncorrosive adhesion over the long term. NHS ester groups, which form covalent bonds with the skin, can be further coupled to the bio-adhesive layer to enhance the couplant’s adhesion on the skin in wet environments such as sweating or soaking in water. The total thickness of the elastomer membrane and the bio-adhesive layer may be selected to be less than % of the acoustic wavelength, thus unaffecting the acoustic transmissivity of the couplant. In a nonlimiting example, the transmissivity may be selected to be >97% relative to degassed water over 48 hours.

10151] When a rigid device is adhered to the skin via the couplant, the interfacial toughness may exceed 100 Jm’² over the long term both in air and under water, maintaining robust adhesion of the device on the skin for any couplant thickness.

(0152] Conventional epidermal devices design includes making the devices thin and stretchable for conformal attachment on the skin. The thin stretchable form factor often compromises the devices’ reliability and functionality yet cannot always guarantee the devices’ stable adhesion on the skin. In some configurations, an ultrasound system in accordance with the present disclosure may be used to robustly adhere thin rigid devices on the skin via a soft, tough, and bio-adhesive coupling layer, which effectively transmits acoustic waves, insulates the devices from skin deformation, and maintains long-term robust and comfortable adhesion on the skin. The bio-adhesive coupling layers may enable electrical, optical, and chemical interfacing with the skin.

[0153] In a non-limiting example couplant formation, chemicals for synthesizing a tough hydrogel may include acrylamide (AAm), acetic acid (AAc), a-ketoglutaric acid, chitosan (high molecular weight), N, N'-methylenebisacrylamide (MBA), calcium chloride (CaC12), K-Carrageenan and the like. The chemicals for synthesizing the elastomer membrane may include hydrophilic polyurethane (Advan Source biomaterials), hydrophobic polyurethane (Advan Source biomaterials), acrylic acid (AAc), benzophenone, a-ketoglutaric acid, l-ethyl-3 -(3 -dimethylaminopropyl) carbodiimide (EDC), and N-hydroxysuccinimide (NHS). The chemicals for synthesizing the bioadhesive layer may include acrylic acid (AAc), acrylic acid N-acryloxysuccinimide (AAc-NHS ester), 2-hydroxy-4'-(2-hydroxyethoxy)-2-methylpropiophenone (Irgacure 2959), polycaprolactone polyurethane, 2-ethylhexyl acrylate (EHA), polyethylene glycol) diacrylate (average Mn 560, PEGDA), ethanol, l-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC), and N- hydroxysuccinimide (NHS).

[0154] In a non-limiting example, to prepare a PAAm-Chitosan tough hydrogel a 3% (w/w) high molecular weight chitosan, 12% (w/w) acrylamide, 0.15% (w/w) N, N'-methylenebisacrylamide, 0.3% (w/w) a-ketoglutaric acid may be disolved in 1% acetic acid solution. The mixture may be centrifuged, such as at 6000 rpm, and poured into a mold, such as a glasses mold with a transparent cover. The solution may then be cured in ultraviolet light (UV) chamber, such as at 364 nm, 10 W power, for 60 min. The resulting hydrogels may then be immersed in IM CaC12 solution for a period of time, such as 24 hours, to reach an equilibrium state.

10155] To prepare a non-limiting example PU elastomer membrane, 30% poly (acrylic acid) may be grafted onto hydrophilic PU by UV-initiated reaction, such as at 364 nm, 10 W power in the presence of 1.5 w/v% benzophenone, and 0.3 w/v% a-ketoglutaric acid for a period of time, such as 120 min. After the reaction, the mixture solution may be dialyzed (Cutoff Mn 3000 Da) against ethanol for a period of time, such as 3 days, and water for 3 days to obtain a pure adhesive PU. The purified adhesive PU may be washed with water and fully dried. In a clean glass, a thin layer of hydrophobic PU (30% w/w) may be spin-coated at 1500 rpm on the glass. After the hydrophobic PU is fully dried, a thin layer of hydrophilic PU (30% w/w) may be spin-coated at 1500 rpm on the hydrophobic PU layer. A thin layer of adhesive PU with EDC and NHS (30% w/w adhesive PU, 5% w/w EDC, and 5% w/w NHS) may be spin-coated at 1500 rpm on the hydrophilic PU layer. After the fdm is fully dried in airflow, such as for 4 hours, the PU elastomer membrane may be obtained.

[0156] To prepare a non-limiting example bio-adhesive layer, 0.5% (w/w) Irgacure 2959, 0.05% (w/w) PEGDA, 12% (w/w) AAc, 35% (w/w) EHA may be dissolved in nitrogen-purged ethanol. The mixture may then be poured on a glass mold with spacers. The adhesive fdm may be obtained after curing the mixture under ultraviolet light (UV), such as 364 nm, 10 W power, for a period of time, such as 40 min.

10157] For forming a hydrogel-elastomer hybrid, the hydrogel may be cut into a desired shape and size, and then the adhesive PU side of the elastomer membrane may be adhered onto the gel with a gentle press to avoid any bubbles. The hydrophobic side of the PU may be adhered onto the bioadhesive layer. The edge of the elastomer and bio-adhesive encapsulation may be cut and sealed by laser sintering. The outer layer of the elastomer membrane may be the bio-adhesive layer, which can adhere to the skin and device firmly after a gentle press.

[0158] The backing layer for the BAUS probe can give mechanical support to the transducers inside the probe that will generate high-frequency vibrations or ARFI. The backing layer also may be selected to have strong attenuation of the ultrasound wave to effectively shorten the pulse duration and thus increase the imaging and elastography resolution. The backing design may be optimized by constructing transducers with backing layers of different compositions of epoxy (EPO-TEK 301) with tungsten powder (particle size 1 um, Sigma-Aldrich), and performing a comparison. In a non-limiting example, a composition of 8: 1 (epoxy to tungsten powder, weight ratio) provided optimized acoustic properties. A BAUS probe with an optimized backing layer may provide for an increase in signal quality, including shorter pulse length and higher signal-to-noise ratio.

[0159] In a non-limiting example, the matching layer for the BAUS probe may be constructed by doping epoxy (EPO-TEK 301) with cerium dioxide nanoparticles (such as with a particle size of 10 nm), in ethanol (cerium dioxide weight ratio in epoxy from 0% to 80%). The matching layer may provide the required acoustic impedance gradient, which may provide for the acoustic energy from the transducer to smoothly penetrate into the body tissue and for the reflected acoustic waves (the returning echo) to smoothly return to the transducer for imaging or elastography.

[0160] As used herein in the context of computer implementation, unless otherwise specified or limited, the terms "component," "system," "module," "controller," "framework," and the like are intended to encompass part or all of computer-related systems that include hardware, software, a combination of hardware and software, or software in execution. For example, a component may be, but is not limited to being, a processor device, a process being executed (or executable) by a processor device, an object, an executable, a thread of execution, a computer program, or a computer. By way of illustration, both an application running on a computer and the computer can be a component. One or more components (or system, module, and so on) may reside within a process or thread of execution, may be localized on one computer, may be distributed between two or more computers or other processor devices, or may be included within another component (or system, module, and so on).

[0161 ] In some implementations, devices or systems disclosed herein can be utilized or installed using methods embodying aspects of the disclosure. Correspondingly, description herein of particular features, capabilities, or intended purposes of a device or system is generally intended to inherently include disclosure of a method of using such features for the intended purposes, a method of implementing such capabilities, and a method of installing disclosed (or otherwise known) components to support these purposes or capabilities. Similarly, unless otherwise indicated or limited, discussion herein of any method of manufacturing or using a particular device or system, including installing the device or system, is intended to inherently include disclosure, as embodiments of the disclosure, of the utilized features and implemented capabilities of such device or system.

[0162] As used herein, the phrase "at least one of A, B, and C" means at least one of A, at least one of B, and/or at least one of C, or any one of A, B, or C or combination of A, B, or C. A, B, and C are elements of a list, and A, B, and C may be anything contained in the Specification.

[01.631 The present disclosure has described one or more preferred embodiments, and it should be appreciated that many equivalents, alternatives, variations, and modifications, aside from those expressly stated, are possible and within the scope of the invention.

Claims

WHAT IS CLAIMED IS:

1 . A method of manufacturing a wearable ultrasound probe, the method including the steps of forming a plurality of slits in a layer of piezoelectric material, each of the plurality of slits having a kerf size less than 40 pm and a center-to-center separation between 150 pm and 250 pm, wherein the slits define edges of a plurality of ultrasound transducer elements formed in the piezoelectric material, the ultrasound transducer elements being configured to produce an acoustic radiation force impulse (ARFI) in a tissue; adhering the layer of piezoelectric material to a matching layer; and implanting a bio-adhesive layer onto the matching layer, the bio-adhesive layer being configured to adhere to skin of a subject.

2. The method of claim 1, wherein the slits are formed in the piezoelectric material using a scratch-dicing method.

3. The method of claim 2, wherein the scratch-dicing method is performed using a dicing saw having a diamond blade with a thickness smaller than the kerf size.

4. The method of claim 1, wherein the layer of piezoelectric material has a thickness between 100 pm and 200 pm.

5. The method of claim 1, wherein the layer of piezoelectric material has a width-to-thickness ratio less than 0.7.

6. The method of claim 1, wherein the plurality of ultrasound transducer elements comprises at least 100 ultrasound transducer elements.

7. The method of claim 1, wherein the bio-adhesive layer comprises a water-based tough hydrogel encapsulated by an elastomer membrane.

8. The method of claim 1, wherein the piezoelectric material comprises lead zirconate titanate (PZT) between an electrode layer disposed on both sides of the PZT.

9. The method of claim 1, further comprising bonding the piezoelectric material to a flexible printed circuit board substrate.

10. The method of claim 1, further comprising filling the slits with epoxy.

11. A wearable ultrasound probe configured to measure shear wave elastography, the probe comprising: an ultrasound array having a plurality of transducer elements, each transducer element comprising a piezoelectric material; and a bio-adhesive layer coupled to the plurality of transducer elements, wherein the bio-adhesive layer is configured to adhere to skin of a subject to position the ultrasound array to acquire shear wave elastography data from the subject; wherein one or more of the plurality of transducer elements is configured to produce an acoustic radiation force impulse (ARFI) that excites a shear wave in a tissue of the subject; and wherein one or more of the plurality of transducer elements is configured to measure shear wave elastography data generated in response to the shear wave.

12. The wearable ultrasound probe of claim 11, wherein the plurality of transducer elements are defined by slits formed in the piezoelectric material, the slits having a kerf size less than 40 pm, and plurality of transducer elements having a center-to-center separation between 150 pm and 250 pm.

13. The wearable ultrasound probe of claim 11, further comprising a matching layer disposed between the ultrasound array and the bio-adhesive layer, the matching layer comprising at least one of silver epoxy or parylene.

14. The wearable ultrasound probe of claim 11, wherein each of the transducer elements is arranged in a linear array.

15. The wearable ultrasound probe of claim 11 , wherein the piezoelectric material is lead zirconate titanate (PZT).

16. The wearable ultrasound probe of claim 15, wherein each of the plurality of transducer elements is formed in the PZT by scratch-dicing the PZT.

17. The wearable ultrasound probe of claim 11, further comprising a processor configured to receive the particle displacement data measured by the transducer elements and generate tissue stiffness measurements based on the particle displacement data.

18. The wearable ultrasound probe of claim 17, wherein the tissue stiffness measurements characterize a stiffness of a tissue 3-5 cm under the skin of the subject.

19. The wearable ultrasound probe of claim 11, wherein the bio-adhesive layer comprises a hydrogel.

20. The wearable ultrasound probe of claim 11, wherein the bio-adhesive layer is configured to be continuously wearable for more than 4 hours.

21. The wearable ultrasound probe of claim 11, wherein the ultrasound array is less than 40 mm in length in an azimuth direction.

22. The wearable ultrasound probe of claim 11, wherein the ultrasound array is less than 26 mm in length in an azimuth direction.

23. The wearable ultrasound probe of claim 11 , wherein the ultrasound array produces the acoustic radiation force using a center frequency of 2-8 MHz.

24. The wearable ultrasound probe of claim 11, wherein the plurality of transducer elements includes at least 100 elements.

25. The wearable ultrasound probe of claim 11, wherein the one or more of the plurality of transducer elements configured to produce an ARFI includes at least 32 transducer elements.

26. The wearable ultrasound probe of claim 11, wherein each of the plurality of transducer elements has a width-to-thickness ratio <0.7.

27. The wearable ultrasound probe of claim 11, further comprising a flexible printed circuit board disposed over the ultrasound array.

28. The wearable ultrasound probe of claim 27, further comprising a backing layer disposed over the flexible printed circuit board.

29. A method for measuring stiffness of a tissue, the method comprising the steps of: a) placing a wearable ultrasound probe on skin of a subject, the wearable ultrasound probe comprising a bio-adhesive configured to adhere the probe onto the skin of the subject and a plurality of ultrasound transducers configured to produce an ARFI in a subcutaneous tissue of the subject; b) using one or more of the plurality of ultrasound transducers to produce an ARFI to generate a shear wave in a tissue of the subject; c) using one or more of the plurality of ultrasound transducers to measure particle displacement data that characterize particle displacement in the tissue generated in response to the shear wave; and d) using a processor to determine a stiffness of the tissue based on the particle displacement data.

30. The method of claim 29, wherein the wearable ultrasound probe remains on the skin of the subject for at least 4 hours, and wherein steps b-d are repeated a plurality of times during the at least 4 hours.