Ultrasound examination has become an essential tool in clinical medicine for a broad spectrum of applications¹. Compared to other medical imaging technologies such as computed tomography (CT) scan, magnetic resonance imaging (MRI), and X-rays, ultrasound stands out for its portability, relatively low cost and greater accessibility². Modern electronics, as well as advances in fabrication and signal processing, have significantly reduced the cost and size of ultrasonic devices. This progress has led to the development of hand-held ultrasonic probes and spurred the adoption of point-of-care ultrasound³. Despite these advancements, ultrasound probes are still too bulky for continuous, skin-mounted monitoring as wearable devices outside of clinical settings, similar to glucose monitors. Moreover, the expertise required for manual operation of these devices by highly trained clinicians further complicates their use in continuous monitoring. Therefore, developing wireless, battery-free, and wearable ultrasonic tools to enable operator-free ultrasound represents a critical step toward integrating continuous diagnostic capabilities into daily clinical care and patient lifestyles⁴.

To address these issues, wearable ultrasound has emerged as a rapidly growing field of interest to both academic and clinical researchers^5,6. Despite being in the early stages of development, wearable ultrasonic transducers have been reported to be used for various applications such as blood pressure monitoring⁷, deep-tissue hemodynamics monitoring⁸, cardiac imaging⁹, elastography¹⁰, long-term organ imaging¹¹, continuous Doppler¹² and bladder volume monitoring^13,14. Most of these reports focused on development of small footprint and conformal ultrasonic transducers, utilizing benchtop equipment for transducer excitation and data processing. Recently, two fully integrated, wireless and wearable ultrasonic patches were demonstrated, including our prior work^14,15. These reports highlight two main challenges for fully wearable ultrasonic patches. Firstly, the electronic circuitry required for wireless operation can occupy several tens of square centimeters, preventing truly conformal, skin-integrated designs^14,15. Secondly, batteries require frequent replacement, hinder long term continuous operation, and depend on external recharging¹⁶ (Supplementary Note 1). The battery issue is also a significant concern for wearable ultrasonic devices since the technology requires high sample rates and processing speeds. For instance, a battery with power consumption of hundreds of milliwatts would require several square centimeters to sustain a few hours of operation. Looking ahead, resolving the size and battery barriers is essential to create miniaturized, electronics-free ultrasonic devices that can be integrated as epidermal patches or contact lenses, enabling new avenues for continuous health monitoring¹⁷, such as continuous blood pressure or bladder monitors, and ultrasonic lens for ophthalmologic ultrasonography. Therefore, similar to bioelectronic implants¹⁸, the realization of fully wearable, battery-free ultrasonic devices capable of further miniaturization remains unachieved, creating an unmet need for a solution that eliminates bulky electronics and power sources (Extended Data Fig. 1).

Here, we introduce ultrasonic tags (US tags) for ultra-compact and electronics-free wearable ultrasound. US tags integrate wireless inductive coupling with ultrasonic transducers in a soft and wearable format for the first time, eliminating the need for bulky batteries and centimeter-scale rigid electronic components. The tags consist of an ultrasonic transducer and a mm-sized coil; and can be carved into both a conformal patch form for epidermal applications and a contact lens form for ocular applications. When wirelessly coupled to an external transmit / receive (T/R) coil, the US tags are able to perform ultrasonic measurements. We demonstrate the potential of this technology for continuous axial eye length monitoring in myopia management, bladder volume assessment in lower urinary tract dysfunctions, and non-invasive blood pressure tracking in cardiovascular care, paving the way for a broad range of future clinical applications.

Design concept of the US tags. Building on the need for fully wearable, battery-free ultrasound systems, the US tags are conceived as distributed sensory nodes at various measurement sites that enable ultrasonic measurements on-demand at multiple sites (Fig. 1a). By separating the transducers from the control electronics and the battery, the ultra-compact US tags can perform wireless ultrasonic measurements at various body locations. Through strategic choices of the placement sites and wireless inductive coupling to the external T/R coil, the US tags enable numerous applications, including the monitoring of axial eye length, bladder volume and blood pressure. Unlike conventional ultrasonic devices, which would require dedicated electronic circuitry for each application, such a wireless measurement scheme streamlines the measurement process, enabling a single external device to interrogate multiple on-body US tags concurrently, reducing cost and complexity.

At its core, the US tags contain an ultrasonic transducer connected to a coil on a single layer printed circuit board (PCB) (Supplementary Fig. 1). The overall footprint of a tag can be further miniaturized by using multi-layer PCBs to implement the transducer coil, and by using piezocomposites to eliminate the need for matching and backing layers in demanding applications. The structure is encapsulated with a biocompatible soft silicone elastomer to conform the human skin. To meet the specific requirements of various diagnostic applications, we design the US tags to cover a range of diagnostic ultrasound frequencies (2, 5 and 10 MHz) balancing penetration depth and resolution to address diverse clinical needs. In addition, to demonstrate different use-cases, transducers with and without matching and backing layers were developed. With these operating frequencies, thicknesses of the US tags range from 300 µm to 1 mm whereas the diameter of the tags are between 1 and 2 cm.

As a step towards a fully portable external device in the future, we also demonstrated the use of the US tags with an evaluation board of an off-the-shelf pulser integrated circuit (Supplementary Fig. 2). By programming the microcontroller unit of the board specifically for the targeted applications, we could excite each US tag with their respective T/R coils from a single external device just by changing the program of the evaluation board.

Characterization of the US Tags. The wireless excitation of the US tags and subsequent wireless reception of the echo signals requires an external T/R coil in close proximity of the US tags (Fig. 2a). By matching the operating frequency of the T/R coil with the US tag, maximum power transfer is achieved, which enables wireless ultrasonic measurements (Supplementary Video 1). Received echo signals are then processed specifically for each application.

Figure 2b displays the simplified equivalent circuit model (ECM) of the wireless US tag system, which guides the design and ensures optimal power transfer and frequency matching. We constitute the US tag ECM by using Butterworth-Van-Dyke (BVD) model to describe the electrical parameters of the PZT based ultrasonic transducer and combine it with the transducer coil inductance L₁ and resistance R₁. When the external T/R coil with inductance L₂ and resistance R₂, which is tuned and matched for the desired frequency using capacitors C_T and C_M, is in close proximity of the US tag, wireless inductive coupling of the ultrasonic signals is achieved.

After ECM based extraction of the required dimensions, the ultra-compact design of the US tag is achieved by advanced laser micromachining techniques to achieve 100 µm line gaps and line widths in transducer coil fabrication (Fig. 2c, Supplementary Fig. 3). Although the parameters vary between the US tags for each application, we develop a generalized design and characterization workflow for the realization of the US tags in three different frequencies. Initially, we extract the electrical parameters (R_m, C_m, L_m and C_o) of the ultrasonic transducer by BVM model fitting (Fig. 2d). For the bladder volume monitoring patch, the resonance is around 2 MHz. We use the extracted parameters as inputs to the model described in Fig. 2b and derive the required L₁ and R₁ values of the transducer coil to match the operating frequency of the US tag with that of the PZT. Figure 2e shows the measured L₁ and R₁ values of a transducer coil that meets the given requirement. This is further proven by the return loss measurements of the US tag with a single turn PCB coil, highlighting the expected frequency of the US tag based on the model and the measured S₁₁ value (Fig. 2g, Supplementary Video 2, Supplementary Figs. 4 and 5).

Figure 2f shows the Laser Doppler Vibrometer (LDV) measurements of the US tag, wirelessly driven by an external T/R coil (Supplementary Fig. 6). Measured surface displacements with respect to excitation frequencies highlight that the maximum displacement of transducer occurs around 2 MHz. Similarly, we characterized the US tags underwater. The pulse-echo response of the wirelessly excited US tag reveals two peaks, highlighting resonance behavior of the US transducer as well as the resonance behavior between the transducer coil and the US transducer (Supplementary Figs. 7 and 8). This is further supported by the underwater frequency sweep of the US tag by a hydrophone, displaying a single peak in agreement with the LDV measurements (Supplementary Fig. 9). To evaluate the underwater pressure field generated by the US tags, we conducted hydrophone measurements in longitudinal and transverse orientations, and observed no discrepancies as compared to conventional wired US transducers (Fig. 2j and 2k).

Bladder Volume Monitoring Epidermal US Patch for Lower Urinary Tract Dysfunctions. Since bladder volume assessment is critical for diagnosing and managing lower urinary tract dysfunctions, we developed an epidermal US patch that can measure diameter of bladder^14,23. Although ultrasound imaging is the most common method to estimate the volume of the bladder non-invasively, it requires clinicians and bulky ultrasound scanners. To enable out-of-clinic bladder volume monitoring, numerous solutions have focused on anterior-posterior wall distance measurements of the bladder as a surrogate measure to the actual bladder volume. Nevertheless, proposed solutions are still bulky, non-conformal and depend on batteries to perform continuous measurements^24–27.

In this work, we integrate the US tags in an epidermal US patch format for bladder volume monitoring by anterior-posterior distance (A-P distance) measurements. With the elimination of the electronics and the battery, the ultra-compact US tag based epidermal patch provides A-P distance measurements in dramatically shrunk form factors.

The patch was characterized by an in vitro setup as shown in Fig. 3a. Three round-bottom flasks (100, 250 and 500 mL) with diameters 6, 8.5 and 11 cm were used to resemble the bladder. Emitted ultrasound waves by the wirelessly excited US tag reflect from the front and back wall of the flasks. After aligning the US tag with the flasks where maximum echo signal is recorded, we can calculate the diameter of the flask from time-of-flight measurements. As seen in Fig. 3b, when used with an alignment platform, the diameter measurements by the wireless US tags yielded a mean relative error of less than 1%, demonstrating high accuracy and reliability comparable to conventional systems.

The performance of the epidermal US patch was further demonstrated in vivo on a healthy volunteer (Supplementary Video 3). As previously mentioned, the echo signals from the anterior and posterior walls of the bladder can act as a surrogate measure for the actual bladder volume. Similar to the in vitro characterization setup, the ultrasound waves emitted by the US tag reflect from various interfaces within the body. Since the bladder is filled with urine, the most distinct signals come from the anterior and posterior walls of the bladder. As urine accumulates, the bladder expands, and it contracts again when urine is voided. The varying A-P distance is a result of this act of continuous expansion and shrinking of the bladder and can be monitored by the echo signals coming from the A-P walls as seen in Fig. 3c. In addition, as can be seen in thermal image in Fig. 3d, even after 5 minutes of continuous operation, the temperature of the not-capsulated transducer coil remains below 30°C.

The results of the wireless US tag-based A-P distance measurements in comparison to ultrasound imaging-based (Fig. 3f, Supplementary Figs. 10 and 11) volume calculations are presented in Fig. 3e. A similar trend between the results of two modalities supports the findings of previous works on the use of A-P distance and highlights the capability of the US tag based epidermal patch to conduct wireless and battery-less A-P distance measurements.

Axial Eye Length Monitoring US Contact Lens for Myopia Management. Numerous medical specialties have benefited from the breakthrough wearable ultrasound technology, but due to the challenges associated with ocular implementations, ophthalmology has not been one of them. Ultrasound is an important tool used by ophthalmologists to assess the anatomy of the eye when optical techniques are unfit, but the technique has been limited to the use of experienced professionals due to the necessity of manual operation and rigid ultrasonic probes²⁸. For instance, axial eye length is a basic anatomical parameter which can be measured by ocular ultrasound, and its monitoring is a critical tool for ocular health due to its relationship with various conditions ranging from myopia to congenital glaucoma and cataract^29,30. This is especially true for myopia, a condition being considered as a global epidemic due to the alarming rates within last decades^31,32. Monitoring and control of axial eye length is regarded as the primary tool for prevention³³, and an ultrasonic device that enables out-of-clinic ocular biometry could be a critical tool for myopia management.

For this reason, we adapt the US tags in a soft contact lens format, presenting the first demonstration of out-of-clinic ocular ultrasound using a fully wearable, battery-free device. As shown in Fig. 4a, the US tag is encapsulated in PDMS inside a 3D-printed mold with a matching radius curvature with the eye phantom. The eye phantom is made up of echogenic layers that resemble the structures of the eye and was used for the in vitro demonstration of the US contact lens as seen in Fig. 4b. When wirelessly excited, the US contact lens successfully captured and transmitted the echo signals from the eye phantom to the external T/R coil for post processing. In addition to the eye phantom, the in vitro performance of the US contact lens was characterized using the setup shown in Fig. 3a. To resemble the eye, two round-bottom flasks with 25- and 50-mL volumes and diameters of 4 and 5 cm were used. As seen in Fig. 4c, when used with an alignment platform, the diameter measurements by the wireless US tags yielded a mean relative error of less than 2%.

The capabilities of the US contact lens were further demonstrated by an ex vivo setup as illustrated in Fig. 4d. In order to artificially cause elongation of the eye, we utilized a syringe pump to fill two enucleated eyes with a certain amount of deionized water. The increase in the pressure due to the injected water caused elongation of the eye. During the injection and elongation, the eyes were monitored with the US contact lens. As shown in Fig. 4e, the US contact lens was able to capture the elongation of the eye through the shifting echo signals (Fig. 4f). With the injection of the water, the structures causing the echo signals move further away from the US contact lens, which was placed on the surface of the eyes. This results in a delay in the echo signal peaks, which carries information regarding the amount of elongation.

In addition to the measurement capabilities, device related factors like stability and biocompatibility are important parameters for the validation. Specifically for the contact lenses, the heating of the coil and the stable performance in eye mimicking medium were chosen for further investigation. As shown in the thermal image (FLIR E96) in Fig. 4g, the transducer coil showed negligible increase in temperature during operation. In addition, there was no shift in the resonance frequency of the US tag when immersed in artificial tear since the whole US tag is fully encapsulated.

Blood Pressure Monitoring Epidermal US Patch for Cardiovascular Care. With the wireless coupling between the control electronics and the US tags, the external electronics can be customized to post-process the received signals to enable more advanced analysis of the time-of-flight measurements. For instance, wearable ultrasound-based measurements on deep vasculature have attracted great interest to enable blood pressure (BP) monitoring^7,19–22. However, the demonstrated transducers require wired connections to benchtop equipment to perform the measurements and process the signals. In this report, we demonstrate that the US tags in epidermal US patch format can eliminate the wired connections from the previously demonstrated BP monitoring wearable ultrasonic devices.

Unlike bladder volume monitoring requirements, which involve deep tissue pulse-echo measurements with less demanding resolution requirements, blood pressure monitoring US tags require high resolution to be able to resolve artery wall movements. Therefore, ultrasonic transducers for BP monitoring were designed with matching and backing layers and encapsulates in PDMS (Fig. 5a). The in vitro characterization setup of the US tag is shown in Fig. 5b. We used a programmable pulsatile pump and a silicone-based vessel to simulate healthy circulation and monitored the flow with both a commercial wired transducer and a US tag. This direct comparison validates the US tag’s ability to provide accurate hemodynamic readings without traditional bulky equipment. The pulsatile nature of the displacement of the walls of the silicone vessel were recorded for both, which were processed and converted in pressure values (Supplementary Note 2). The agreement between the two measurements for an extended time duration and the input command of the pump piston displacement can be seen in Fig. 5e. All phases of a cardiac cycle were successfully captured and the distinct points of the systolic, dicrotic and diastolic phases were clearly visible.

For proof-of-concept, in vivo blood pressure monitoring capability of the device was demonstrated on a healthy volunteer (Supplementary Video 4). The US tag in epidermal US patch format was manually positioned by the volunteer according to the ultrasound images taken by a commercial probe (Fig. 5c). The signal flow of the blood pressure measurement can be seen in Fig. 5d. When the T/R coil is excited, only the transmit waveform is observed on the oscilloscope screen without any echoes. As the T/R coil is brought in close proximity of the US tag, distinct echoes can be seen. With the appearance of these echo signals, the algorithm initiates a calibration phase which detects the echoes corresponding to the artery wall. These echo movements are then translated into artery diameter and blood pressure readings. Representative BP waveforms from a commercial ultrasonic transducer, the ultrasonic transducer of the US tag in wired form and wireless US tags are shown in Fig. 5f. The wirelessly coupled US tag successfully captured the major changes in the cardiac cycle. When the coils are decoupled, the echo signals from the artery walls disappear as shown in Fig. 5g and Supplementary Video 4. In accordance with the reported methods, the pressure values from a commercial cuff measurement were used for the calibration, and systolic and diastolic pressure values from the US tags were taken as the mean values recorded during a 5-second interval. As seen in Fig. 5h, the US tags accurately captured the heart rate and the blood pressure values when compared with the values obtained from the cuff measurements.

Thanks to the advances in soft materials, fabrication techniques and sensing strategies, wearable medical devices are revolutionizing the healthcare system by enabling continuous and quantitative assessment of various biomarkers^34,35. Nevertheless, clinical adaptation of such breakthrough devices is still at its infancy due to several technological barriers³⁶, including size constraints, energy management, and integration challenges—areas directly addressed by our battery-free, ultra-compact US tags. Firstly, albeit the sensor miniaturization efforts, the use of centimeter-scale rigid electronic components and batteries for physiological monitoring and wireless data transmission impedes the transition from the large and bulky systems³⁷. In addition, the power demand of the wearables increases as the sensors get more sophisticated and multiplexed³⁸.

Inspired by the researchers from various fields who utilized wireless signal transmission between two coils to overcome these challenges^39–45, we introduce the US tags to tackle the major problems of the wearable ultrasound technology. Firstly, the US tags drastically reduce the footprint of the device by removing the electronics and the battery. Miniaturized form factor not only improves the wearability immensely for existing applications but also enables new applications like out-of-clinic ophthalmologic ultrasound, for the first time. Moreover, existing applications with wired approaches require dedicated electronics and battery for each application. With the separation of the control electronics and the ultrasonic transducers, the users can utilize a single device to conduct measurements with all of the worn US tags.

Future work could focus on optimization of the transducers of the US tag for specific applications. For instance, lead-free alternatives could be implemented instead of PZT for environmental concerns, and piezoelectric micromachined ultrasonic transducers could be adapted for further miniaturization potential. In addition, the current readout scheme can be further miniaturized and integrated into truly portable devices, potentially interfacing with smartphones or tablets to enable rapid, real-time diagnostics and telemedicine applications. As the first step, we utilized an evaluation board of an off-the-shelf pulser integrated circuit and customized three programs of its microcontroller to demonstrate all three applications from a single device. Finally, for the applications where long-term wearability is critical, skin couplant layer of the US tag could be optimized to improve breathability and resistance to perspiration.

Fabrication US tags and T/R System. Wrap-around electroded PZT discs and bulk PZT plates with silver electrodes were purchased from American Piezo (APC − 850) and CTS Ferroperm (Pz27). Bulk PZT plates were cut into desired dimensions by dicing (DAD3350, DISCO Corporation). To enable access to both electrodes of the diced PZTs from the same plane, a portion of the electrode on one face of the PZTs was laser ablated (500 kHz pulse repetition frequency with a divider of 3, 1.5 W power and 1500 mm s^− 1 laser cutting speed, R4, LPKF Laser & Electronics). The isolated portion of the silver electrode on laser ablated face was electrically connected to the electrode on the opposite face by low temperature silver paste (DuPont, PE827).

200 µm thick flexible laminate with 18 µm copper layer (LPKF, 119818) was patterned with LPKF U4 to fabricate the flexible coils of the tags, such that the operation frequency of the US tag coincides to the resonance frequency of the PZT transducer. Prepared PZTs were manually aligned with the pads of flexible coils and reflowed with low-temperature solder paste (TS391LT, Chip Quik). Finally, degassed silicone (PDMS) elastomer mixture (Dow Corning, Sylgard 184 10:1) was used as the encapsulation material with dedicated molds for epidermal patch and contact lens formats and cured in an oven at 40°C overnight (Supplementary Figs. 12 and 13). Lastly, a standard 1.5 mm thick FR4 board with 18 µm copper layer (LPKF) was patterned with LPKF U4 to fabricate the T/R coils of the tags.

For blood pressure monitoring application, matching and backing layers were used to improve the device performance (Supplementary Fig. 14). First, lead wires were soldered on wrap-around PZTs using low temperature solder wire (SMDIN52SN48, Chip Quik). The soldered PZT was placed inside a silicone mold with a cavity and was transition fitted into the cavity. A few layers of a PTFE double-sided tape was then laser patterned and adhered on to the ceramic as the spacer for matching layer screen printing step. At the same time, epoxy resins filled with various inorganic particles were prepared and characterized to be used as the matching and backing layers (Supplementary Fig. 15). For the matching layer, 60% alumina nano particle (290 nm, Nanografi) loaded epoxy resin (EPOTEK 301-2 FL, 1:0.35) was prepared by speed mixing for 15 minutes and then degassing for 20 minutes. Prepared composite resin was screen printed onto the ceramic, and the assembly was cured in an oven at 80°C for 8 hours. Similarly, after flipping the ceramic/matching layer assembly, the backing layer consisting of tungsten microparticles (1–10 µm, Nanografi) and highly attenuative epoxy was prepared, screen-printed and cured. The matching layer was than lapped to the desired thickness with the help of conductance measurements¹⁴ (Supplementary Fig. 16).

Characterization of US tags and T/R System. An impedance analyzer (MFIA, Zurich Instruments) was used to determine the electrical characteristics of the transducer. Measured impedance values were fitted to Butterworth-van-Dyke model for subsequent coil design.

The coil parameters were chosen such that the US tag operates at the resonance frequency of the PZT transducer. Impedance values of the fabricated coil was determined by the impedance analyzer. After the fabrication, S₁₁ parameters of the US tag was wirelessly measured using a single turn planar spiral coil connected to a vector network analyzer (VNA) (SVA1000X, Siglent).

Impedance matching and tuning capacitors were used to maximize the power transfer and tune the frequency to the frequency of the tags. During matching and tuning, the coil was connected to the VNA to determine the S11 parameter.

Surface displacement of the wirelessly excited US tags was characterized with Laser Doppler Vibrometer (LDV) (MSA-600, Polytec). Coils were coaxially aligned and affixed with an inter-coil distance of 5 mm. T/R coil was excited with a 2 V periodic chirp signal generated by the LDV, and surface displacements were recorded for the frequency range of interest.

To evaluate the wireless pulse-echo response, US tags were placed in an acrylic water tank and affixed on the wall with a double-sided tape. The tank was filled with deionized water, and T/R coil was aligned from outside the water tank. An arbitrary waveform generator (AWG) (33521B Waveform Generator, Keysight) was used to excite T/R coil. The echo signal from a stainless-steel reflector was received by an oscilloscope (Picoscope 6000 series) through the use of a diplexer (RDX-6, RITEC Inc.)

Further acoustic characterizations were conducted underwater inside an UMS Research system covered with acoustic absorbers (AptFlex F28, Precision Acoustics). Similar to pulse-echo response, the US tags were affixed inside the water tanks and T/R coils were aligned from outside. The longitudinal and transverse pressure field, and the bandwidth were evaluated with a 0.5 mm needle hydrophone (Precision Acoustics) connected to an oscilloscope (DSOX3024G, Keysight Technologies).

In vitro characterization. Given the shapes of the eye and the bladder, axial eye length monitoring and bladder volume monitoring applications were simplified into a task of diameter measurement of spherical/ellipsoid objects. With this in mind, spherical-shaped round bottom flasks were connected to a motorized stage and submerged in DI water, with US tags affixed inside the water tank. Considering the scale of the eye and bladder, 25 and 50 mL flasks were used for the US contact lens, and 100, 250 and 500 mL flasks were used for the bladder monitoring US tag. With the help of the motorized stage, the flasks were aligned with the US tags to maximize the time between the echoes reflected from the anterior and posterior walls of the flasks. Similar to the procedure above, AWG was used to excite the T/R coil, and echoes were displayed and recorded on the oscilloscope through the use of the diplexer.

In addition to underwater flask experiments, an eye phantom (VK-10222, Gampt mbH) was used to evaluate the performance of the US contact lens with a silicone-based couplant (Silbione 4717 (1:1), Elkem). T/R coil was brought in close proximity of the phantom and the US lens and was excited by pre-programmed ultrasound pulser evaluation board (STEVAL-IME013V1, STMicroelectronics). Received echoes were again displayed and recorded on the oscilloscope.

For the in vitro characterization of the heart rate and blood pressure monitoring US tag, a pulsatile pump (SuperPump, ViVitro Labs) with a silicone rubber vessel was used. The pump was programmed to replicate physiological flow profile with 70 beats per minute, and a custom setup with a water tank containing the pulsating silicone vessel was built. Similar to underwater characterizations, the US tag was aligned with the pulsating vessel and affixed onto a wall of the water tank. T/R coil was aligned with the US tag and affixed onto the other side of the water tank and was excited by the AWG together with an amplifier (F30PV, Pendulum).

In vivo bladder volume measurements. Bladder volume measurements were carried out on a healthy male volunteer with an approved protocol by the Koç University Ethical Committee. Measurements were taken at 10-minute intervals while the US tag was placed on the lower abdomen. For ultrasound imaging, ultrasound probe was manually placed longitudinal and transverse orientation until the maximum dimensions of the bladder was observed on the scanner’s screen and recorded. Length, width and height dimensions of the bladder were then manually determined by caliper placement on the recorded images. Finally, prolate ellipsoid volume method was used to estimate the bladder volume from the recorded US images.

Ex vivo characterization of the US Contact Lens. Two enucleated cow eyes were used to evaluate the axial length monitoring capabilities of the US contact lens. Vitreous chambers of the eyes were cannulated with needles to inject deionized water to cause elongation of the axial eye length due to increased pressure. Syringe pump (SyringeONE, New Era Instruments) was used with an infusion needle to control the amount of injected DI water. At the same time, US contact lens was placed onto the eye-under-experiment. After every 0.5 mL of DI water injection, pulse-echo measurements were conducted with the setup used during in vitro experiments. Received echoes were later processed by using band-pass filtering, time-gain compensation and amplification.

In vivo blood pressure measurements. Blood pressure measurements were carried out on a male volunteer with an approved protocol by the Koç University Ethical Committee. A commercial digital blood pressure monitor was used for conventional blood pressure measurements and the obtained pressure values were used as calibration data for US tag-based measurements.

A blood pressure algorithm was designed to assess the artery diameter and blood pressure in real-time (Extended Data Fig. 2). Initially, blood pressure values were measured by a cuff and used as reference. The received signal is then band-pass filtered to remove noise, followed by an envelope generation via the Hilbert transform to get a one-sided amplitude profile of the signal. Before real-time blood pressure monitoring, the algorithm initiates a calibration process, during which the peaks within the echo signals are identified as candidate regions of interest (ROIs). Given the presence of stable and moving echo signals, originating from various tissue regions, the algorithm specifically identifies arterial echo signals. To distinguish these, a ranking algorithm evaluates echo movements through cross-correlation with consecutive signals, effectively differentiating movements and stable echoes. This ranking isolates signals with the most significant oscillations, enabling the identification of the arterial anterior and posterior wall echoes based on their temporal separation. These anterior and posterior wall ROIs serve as references in the main processing stage. To track the movement of the anterior and posterior walls, new frame ROIs undergo further cross-correlation and time-of-flight values that relate to the artery diameter, and ultimately blood pressure values are calculated using the time-of-flight values of the anterior and posterior wall ROIs. Smoothing is applied at the final step, enhancing the stability and readability of both the blood pressure and artery diameter results.

The authors declare that the data supporting the findings of this study are available within the paper and its supplementary information files. The data belonging to the study will be available in Koç University Institutional Repository (KUIR) as soon as possible. In the meantime, they are available for research purposes from the corresponding author upon reasonable request.