I. Core Research Summary

• Inductively Coupled Passive Battery-Free Architecture

The stent is laser‑processed into a helical conductive circuit that serves as a wireless receiving antenna. At its two ends, aerosol‑printed dual‑capacitor sensors form an LC resonant circuit, eliminating the need for an internal power source. An external antenna remotely detects changes in the resonant frequency, which are then converted to yield real-time intravascular pressure and pressure differentials, enabling indirect estimation of blood flow.

• Aerosol‑Molded Highly Flexible Microstructured Capacitive Sensor

Layered aerosol printing constructs a PI substrate, a silver‑nanoparticle conductive layer, and a PDMS micro‑convex dielectric layer; the microstructured dielectric markedly enhances pressure sensitivity. The sensor withstands an extremely small bending radius of just 0.25 mm, and its structure remains intact under 14 atm of expansion pressure. After 2,500 pressure cycles, its performance remains stable with no drift.

• Ultra-long wireless reading distance

The maximum radial readout distance in air is 5.5 cm, and in a saline‑based blood‑mimicking environment it reaches 3.5 cm—far exceeding that of comparable vascular implantable sensing devices (typically ≤1 cm). Stable signal acquisition is also achievable in subcutaneous tissue.

• Minimally invasive catheter-compatible integrated device

The stent has an initial diameter of 1.5–2 mm and is balloon‑expanded to 3–5 mm, with the sensor conforming to the stent’s inner wall in an ultra‑thin, low‑profile configuration. A minimally invasive approach via the carotid artery was used to deliver the device into the rabbit iliac artery, with no kinking or vascular injury throughout the procedure; the device can remain in place for long-term monitoring after implantation.

• Simultaneous assessment of multiple hemodynamic parameters

The dual-sensor units at both ends of the stent create a pressure-difference detection zone; pressure is measured at a single point, and the pressure difference between two points is used to calculate blood flow velocity. The device provides real-time output of three sets of data—systolic pressure, diastolic pressure, heart rate, and blood flow volume—enabling precise identification of abnormal flow fluctuations caused by vascular stenosis.

II. Core Operating Principle of Aerosol Spraying and Its Unique Process Advantages

(1) Complete Printing Process for Sensing Thin Films

• PMMA sacrificial layer spin-coated onto a glass substrate; plasma activation enhances ink adhesion.
• Pneumatic atomization PI‑insulated substrate ink, with aerosol‑based underlayer printing and curing at 240°C;
• Ultrasonic atomization of silver nanoparticle (AgNP) conductive ink is used to print the upper and lower capacitor electrodes, followed by high-temperature sintering to reduce sheet resistance.
• Precise patterned aerosol‑jet printing of a micro‑convex PDMS dielectric layer to fabricate a microscopic elastic sensing structure.
• The sacrificial layer is removed by acetone dissolution, and the fully integrated flexible sensor is peeled off, transferred, and firmly bonded to the inner wall of the sensing substrate.

(II) Five Core Processes (in comparison with spin coating, blade coating, and vacuum evaporation)

• Direct-write capability of microstructured dielectric layers

A digital printing–based array of PDMS protrusions can be fabricated without mold‑based embossing, enabling one‑step formation of elastic microstructures and substantially enhancing pressure‑sensing sensitivity, whereas conventional coating methods are limited to producing only flat thin films.

• Ultra-thin, low-profile, multi-layer integrated molding

Multilayer conductive and insulating materials are precisely deposited in sequence, with a total sensing thickness of less than 30 μm. This ultra-thin design does not increase the overall outer diameter of the stent and does not impede intravascular blood flow.

• Conformal Deposition on Curved/Non-Planar Substrates

The aerosol‑free mist stream conforms to the curved inner wall of the support, forming a uniform film inside the narrow cylindrical cavity, thereby addressing the challenge of flat‑bed printing being unable to accommodate tubular implant carriers.

• Compatible with wide-viscosity biomedical inks

It is compatible with three vastly different types of inks—PI, silver nanoparticle, and PDMS—and enables the complete fabrication of electrodes, dielectrics, and substrates in a single machine, eliminating the need to switch production lines.

• Digital rapid iteration, no photolithography mask required.

CAD enables free adjustment of electrode dimensions and PDMS micro‑pillar arrays, while sensor prototypes tailored to various tube diameters can be fabricated in just a few hours, significantly shortening the R&D cycle for implantable devices.

III. Five Key Technological Innovations

Innovation 1: A Novel Process for Fabricating Aerosol-Microstructured Capacitive Sensing Films

Innovative dual-sensor differential pressure-based blood flow detection architecture

Innovation: 3‑in‑1 inductor bracket + aerosol sensing integrated design

Innovative, highly tolerant interventional design with flexible sensing technology

Innovation 5: Optimization of Long-Range In Vivo Wireless Sensing Systems

IV. Existing Process Challenges and Corresponding Solutions for Overcoming Them

Challenge 1: The flat PDMS membrane exhibits extremely low pressure sensitivity, resulting in poor signal discrimination.

       Challenge 2: Traditional printed films are thick, increasing the overall outer diameter of the stent and thereby raising the risk of vascular occlusion.

Challenge 3: High-pressure balloon inflation and repeated bending of the vessel can cause the conductive layer to fracture.

Challenge 4: Blood’s high electrical conductivity significantly attenuates RF wireless signals, resulting in an extremely short read range.

Challenge 5: Molded micro‑media require custom molds, making it difficult to quickly adapt to stents of different diameters.

V. Future Diversified Industrial Application Scenarios

1. Intelligent Interventional Stent for Coronary and Peripheral Vessels

2. Long-term monitoring device for intracranial aneurysms

3. Animal Cardiovascular Disease Research Platform

4. Dialysis Vascular Access Monitoring Stent

5. Minimally Invasive Vascular Interventional Diagnosis and Treatment Integrated Device

Figure 1. Overview of the fully implantable wireless vascular electronic system, equipped with printed sensors for wireless hemodynamic monitoring. (A) Schematic of the implantable electronic components. (B) Inductive stent design constructed using conductive gold rings and non-conductive polyimide (PI) connectors; the current path mimics a solenoid structure (left panel), along with a scanning electron microscopy (SEM) image of the stent (right panel). (C) Layered structure of the soft pressure‑sensitive sensor (including a printed dielectric layer, left panel) and a photograph of an index finger simultaneously holding both the flow sensor and the pressure sensor (right panel). AgNP: silver nanoparticles; PDMS: polydimethylsiloxane. (D) Schematic illustrating minimally invasive catheter insertion and balloon expansion of the wireless vascular stent. (E) Initial and expanded states of the sensor‑integrated stent system. (F) The wireless stent system implanted in the right iliac artery of a live rabbit. (G) Schematic of the wireless design and sensing scheme, enabling simultaneous monitoring of pressure, heart rate (HR), and blood flow parameters.

Figure 2. Design, fabrication, and performance characterization of the wireless stent. (A) Schematic illustration of the fabrication steps for the multi-material inductive stent. (B) Layout diagram of the stent structure employing conductive gold rings and non-conductive polyimide (PI) connectors. (C) Scanning electron microscopy image of the stent structure showing the polyimide connectors; a magnified view of the PI connector clearly reveals the separated state of the gold rings. (D) Cross-sectional image of the stent struts, illustrating the layered architecture comprising stainless steel (SS), a gold coating, and a parylene layer. (E) Balloon inflation process of the wireless stent: resistance increases only slightly, while inductance rises markedly; photographs document the dynamic changes during balloon expansion. (F) Trend of increasing stent diameter as balloon pressure is raised. (G) Comparison of the magnitude of the S11 parameter at the resonant frequency between the wireless stent and a copper coil at various distances—when using a single-loop mismatched antenna for wireless readout, the parameter magnitude decays with increasing distance and becomes undetectable below the noise floor. (H) Maximum wireless readout distances achieved by each stent under single-loop antenna conditions; employing a matched antenna significantly enhances the readout range. (I) Wireless spectrum scan results of the S11 parameters for the stent and sensor: larger separation distances result in a reduced attenuation of the resonance signal. (J–L) Measured values of axial stiffness (J), bending stiffness (K), and radial stiffness (L) for the wireless stent, compared with those of commercial stents, a wireless stent without polyimide connectors, and an inductive stent design utilizing stainless steel connectors.

Figure 3. Fabrication and performance characterization of the soft pressure sensor. (A) Schematic cross-section illustrating the layered structure of the sensor, prepared by sequentially jet‑printing polyimide (PI), silver nanoparticles (AgNP), and a dielectric polymer (PDMS) ink. The top and bottom electrodes are printed separately before being assembled into a composite structure. (B) Photograph of the aerosol‑jet nozzle used for printing the sensor. (C) A soft pressure sensor with an interconnect structure mounted on a fingertip. (D) Scanning electron microscope image of the sensor’s bottom electrode, comprising a supporting PI layer, a conductive AgNP layer, and a dielectric PDMS layer. (E) Profile measurements of the interconnect structure (left), the electrode (middle), and the dielectric layer (right). (F) Compared with solid thin films of the same thickness, employing a patterned PDMS layer as the dielectric significantly enhances pressure sensitivity. (G) The sensor’s capacitance variation over time (left) and its response to abrupt large pressure changes (right) both closely match the pressure‑wave response; the device provides instantaneous feedback to pressure steps of 300 mmHg. (H) After 2,500 cycles of pressure ranging from 0 to 1,000 mmHg, the sensor exhibits negligible performance degradation. (I) Sensor response characteristics during balloon inflation, in conjunction with a wireless support structure. (J) Verification that the sensor remains functional under torsional and bending conditions. (K) Pressure‑response behavior under bending: sensitivity remains stable at a bending radius of 1.5 mm, and the sensor retains detectability even beyond a bending radius of 0.25 mm. (L) Comparative analysis between this sensor and existing approaches for pressure sensing under bending. This study demonstrates that, among capacitive pressure sensors, this device achieves the smallest bending radius, ranking second smallest among two distinct sensor types.

Figure 4. Wireless sensing demonstration of pressure, pulse, and flow. (A) Schematic of the wireless sensing system through a guiding catheter. (B) Expanded stent and sensor architecture within an arterial model; the inset shows a cross-section of the low-profile electronic components, while the magnified view illustrates the expanded stent structure and the PI connector. (C) Schematic of wired versus wireless sensing approaches in the arterial model. (D) Amplified plot of sensor capacitance variation and pressure waveforms under pulsatile flow conditions in the arterial model. (E) Summary of wired sensor response characteristics at different flow rates: capacitance increases linearly with pressure, indicating minimal influence from fluid flow. (F) Wireless resonant frequency scan curves at varying pressures: resonant frequency decreases as pressure rises. (G) Wireless pressure-sensing system in the arterial model, along with an amplified view of the pulsatile waveform. (H) Summary of wireless sensing data for mean, peak, and trough pressures during pulsatile flow. (I) Pulse rate measurements under two flow regimes: the pulse rate obtained by the wireless sensor is consistent with that measured by a commercial pressure sensor. (J) A wireless stent integrating dual pressure sensors for blood‑flow monitoring: the two sensors simultaneously track two resonant frequencies, enabling real-time monitoring of the pressure gradient (ΔP). (K) Comparison of wireless blood‑flow monitoring performance: analysis of differences in pressure gradient measurements between the wireless sensor and a commercial sensor. (L) Schematic of wireless signal behavior when the device operates in air versus physiological saline; conductive environments degrade wireless signal quality. (M) Wireless readout distances in air, in physiological saline, and in a saline–tissue mixture. (N) Schematic of the external antenna and the arterial model embedded in meat at maximum readout distance.

Table 1. Performance Comparison of Implantable Wireless Arterial Sensors and Systems

*Graft fixation is performed once the distance has been reduced to 3.0 cm (for blood‑flow monitoring). §Wireless signal strength is rated on a scale of 1 to 5; †Thickness is estimated by summing the stent and sensor thicknesses (in centimeters), though the specific required strength value is not specified; ‡The device is implanted inside the graft; ||The device is implanted using a wired connection; ¶This device measures pulse rate and distinguishes between the presence and absence of blood flow; #Total thickness is calculated by adding up the thicknesses of each reported layer; **The device is wrapped around the arterial surface and secured with sutures to the surrounding tissue; ††Thickness of the sensor in the absence of a stent; ‡This device must be fully implanted, unlike other devices that can operate independently or self‑support; §§Specific distance values are not specified, but the CardioMEMS HF system has already entered the human clinical trial phase.

Figure 5. In vivo studies of transcatheter implantation. (A) Photograph showing a 1.5‑mm‑diameter stent and sensor advanced through a guiding catheter and expanded over a balloon. (B) Photograph of the stent and sensor after expansion. (C) Schematic illustration of catheter-based implantation in a rabbit: the wireless device is advanced along a balloon‑tipped catheter from the common carotid artery to the right iliac artery. (D) Fluoroscopic image demonstrating the target site in the right iliac artery; following delivery of the stent and sensor to the right iliac artery, the balloon is inflated and the catheter is withdrawn. (E) Image of the implanted stent and sensor within the right iliac artery. (F) Wireless frequency‑scan signals from the implanted sensor before and after removal; after three months, only minor changes in the signal are observed. (G) Resonant frequencies before and after balloon inflation are consistent with theoretical predictions. (H) Wireless testing of the implanted device in an arterial model confirms its pressure‑sensing capability.