An implantable pressure sensor and a preparation method and application thereof

Abstract

This invention belongs to the field of biomedical materials and flexible sensing technology, and particularly relates to an implantable pressure sensor, its preparation method, and its application. The pressure sensor has a three-layer parallel plate structure, wherein the electrode layer is a conductive hydrogel prepared from polymerizable zwitterionic monomers, the dielectric layer is a microphase-separated hydrogel formed by copolymerization of acrylamide monomers and acrylate monomers, and the encapsulation layer is a biocompatible polymer fiber membrane. By controlling the material composition, polymerization conditions, and structure, the device achieves good matching with biological soft tissues in terms of water content and tissue adaptability. In terms of functional characteristics, the implantable sensor exhibits advantages such as sensitive pressure response and strong cyclic stability, and can be applied in fields such as tissue engineering, physiological or pathological signal monitoring, neural stimulation modulation, and drug delivery.

Description

Technical Field

[0001] This invention belongs to the field of biomedical tissue engineering materials technology, specifically relating to an implantable pressure sensor, its preparation method, and its application. Background Technology

[0002] Implantable devices have been widely used in tissue engineering, biomonitoring, neuromodulation, and in vivo therapy. To achieve stable and efficient applications, these devices need to maintain good structural coordination and functional coupling with surrounding tissues during operation in complex and dynamic physiological environments. Currently, most implantable devices are fabricated using metal and silicon-based materials. While these materials offer advantages in mechanical strength and processing performance, their physicochemical properties differ significantly from those of biological soft tissues. This difference can easily lead to stress concentration, frictional stimulation, chronic inflammation, and signal distortion at the device-tissue interface, thus limiting the device's performance and lifespan.

[0003] Hydrogels, as soft materials composed of polymer networks and a large number of water molecules, are gradually becoming important candidate materials for implantable devices due to their high water content similar to human soft tissue, controllable mechanical properties, and good biocompatibility. However, existing hydrogels still have the following shortcomings: their mechanical strength, toughness, and fatigue resistance are insufficient, making it difficult to adapt to the complex mechanical environment in vivo, and they are prone to structural damage or performance degradation under long-term loads. Furthermore, they are also prone to excessive swelling, component loss, and weak stability in the body fluid environment. At the same time, it is difficult to synergistically optimize their functional properties (conductivity, tissue adhesion, stimulus responsiveness, etc.) and biocompatibility. In addition, the interfacial bonding stability between hydrogels and device substrates or functional layers is insufficient, and interfacial delamination and functional failure are prone to occur in the body fluid environment and under repeated mechanical action.

[0004] Therefore, it is essential to provide an implantable pressure sensor and its preparation method, which uses molecular structure design, multi-component composite and microstructure regulation to obtain hydrogel materials that are closer to natural biological tissues in terms of water content, mechanical properties, biocompatibility, interface stability and functional adaptability, and to construct implantable devices suitable for complex in vivo environments. Summary of the Invention

[0005] Based on the above background, the purpose of this application is to provide an implantable pressure sensor and its preparation method. Through molecular design and structural optimization, the materials and devices are made closer to natural tissues in terms of biocompatibility, mechanical adaptability and long-term stability, thereby effectively reducing immune response, inflammation and fibrous encapsulation formation, and avoiding a decrease in sensing sensitivity.

[0006] This invention proposes a method for fabricating an implantable pressure sensor, the method comprising the following steps: Step S1: Prepare zwitterionic conductive hydrogels using zwitterionic monomers; Step S2 involves copolymerizing acrylamide monomers and acrylate monomers as two components to prepare a microphase-separated copolymer hydrogel.

[0007] Step S3: The conductive hydrogel obtained in step S1 is used as the upper electrode layer and the lower electrode layer, and the microphase-separated copolymer hydrogel obtained in step S2 is used as the dielectric layer. The layers are assembled to obtain a three-layer structure device. Step S4: An encapsulation layer is provided on the outside of the three-layer structure device to obtain an implantable pressure sensor.

[0008] Specifically, in step S1, a prepolymer solution is prepared by adding a crosslinking agent and an initiator to a polymerizable zwitterionic monomer; after sufficient polymerization to obtain a hydrogel, it is soaked in pure water to remove unreacted chemicals and reagents, thereby obtaining a zwitterionic conductive gel; the polymerizable zwitterionic monomer includes 3-[[2-(methacryloyloxy)ethyl]dimethylammonium]propionate, [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide and 2-methacryloyloxyethyl phosphocholine.

[0009] Specifically, the crosslinking agent is polyethylene glycol diacrylate (PEGDA), and the initiator is 2-hydroxy-2-methyl-1-[4-(2-hydroxyethoxy)phenyl]-1-propanone (Irgacure 2959). The degree of crosslinking of the zwitterionic hydrogel is controlled by adjusting the amount of crosslinking agent, thereby adjusting the mechanical properties and ion conductivity of the hydrogel.

[0010] Hydrogels with different degrees of crosslinking were prepared by adjusting the amount of crosslinking agent to optimize their pressure sensing sensitivity; preferably, the mass ratio of zwitterionic monomer, crosslinking agent and initiator was 2.1:0.025:0.01.

[0011] Specifically, in step S2, when copolymerizing acrylamide monomers and acrylate monomers as two components, a crosslinking agent, initiator and catalyst are added to prepare a prepolymer solution; after polymerization is complete to obtain a hydrogel, it is soaked in pure water to remove unreacted chemicals and reagents to obtain a microphase-separated copolymer hydrogel; the initiator is ammonium persulfate.

[0012] Specifically, the acrylamide monomers mentioned in step S2 include one or more combinations of acrylamide, methacrylamide, N,N-dimethylacrylamide, N-hydroxyethylacrylamide, N-isopropylacrylamide and their salts; by adjusting the ratio of acrylamide monomers to acrylate monomers and the polymerization temperature, the dielectric properties, compressibility and wet adhesion properties of the microphase separation copolymer hydrogel are controlled.

[0013] Specifically, in step S2, the crosslinking agent is polyethylene glycol diacrylate (PEGDA), the initiator is ammonium persulfate (APS), and the catalyst is tetramethylethylenediamine (TEMED); preferably, the ratio of acrylamide monomer to acrylate monomer is 2:1.

[0014] Specifically, in step S4, the encapsulation layer is an electrospun aliphatic polyester material; the aliphatic polyester material includes, but is not limited to, polycaprolactone, polylactic acid, lactic acid-glycolic acid copolymer, and polyglycolic acid.

[0015] The present invention also proposes an implantable pressure sensor, the sensor comprising an electrode layer, a dielectric layer and an encapsulation layer; wherein the upper and lower electrode layers are both zwitterionic conductive hydrogels, the dielectric layer is a microphase-separated copolymer hydrogel, and the encapsulation layer is an aliphatic polyester material.

[0016] The present invention proposes the application of an implantable pressure sensor in physiological signal monitoring, wherein the physiological signals include tissue pressure signals, visceral motion signals, etc.

[0017] Compared with the prior art, the present invention has the following advantages: The pressure sensor prepared in this invention has a three-layer parallel plate structure. The electrode layer is a conductive hydrogel prepared from polymerizable zwitterionic monomers, the dielectric layer is a microphase-separated hydrogel formed by copolymerization of acrylamide monomers and acrylate monomers, and the encapsulation layer is a biocompatible polymer fiber membrane. The tissue-mimicking hydrogel-based implantable pressure sensor prepared in this way has high mechanical compatibility with human soft tissue and excellent biocompatibility, which can significantly reduce interfacial stress and chronic inflammation, and ensure long-term stable operation in vivo. The sensor has high sensitivity and good linear response within the physiological pressure range, strong cyclic stability and recoverability, and strong interlayer bonding and interface stability, which can maintain reliable function in complex body fluid and mechanical environments. The device has good in vivo biodegradability, eliminating the need for secondary surgery for removal, effectively reducing patient trauma and medical costs. Through flexible control of material composition and preparation process, the device performance can be customized, making it widely applicable to multiple biomedical fields such as tissue engineering, physiological signal monitoring, neural stimulation modulation, and drug delivery. Attached Figure Description

[0018] To more clearly illustrate the technical solutions in the embodiments of the present invention, the drawings used in the description of the embodiments or the prior art will be briefly introduced below. Obviously, the drawings described below are only some embodiments of the present invention. For those skilled in the art, other drawings can be obtained based on these drawings without creative effort.

[0019] Figure 1The images show the zwitterionic monomer solution (left) and zwitterionic hydrogel (right) prepared according to the present invention. Figure 2 Comparison of the microphase separation hydrogel of this invention before and after swelling; Figure 3 This is a microstructure diagram of the electrospun fibers prepared in this invention; Figure 4 This is a schematic diagram of the capacitive pressure sensor structure prepared according to the present invention. Figure 5 The attached diagram shows the apparent viscosity of the hydrogel prepared in this invention. Figure 6 The degradation curve of the hydrogel prepared in this invention is shown. Figure 7 The results of cell compatibility experiments on the sensor components prepared in this invention; Figure 8 A schematic diagram illustrating the principle of pressure response testing for this invention; Figure 9 The sensitivity curve of the pressure sensor prepared according to the present invention is shown. Figure 10 This is a stability test diagram of the pressure sensor prepared according to the present invention; Figure 11 The pressure sensor prepared according to the present invention is shown in the gastric peristalsis signal response curve. Detailed Implementation

[0020] In the following description, specific details such as particular system architectures and techniques are set forth for illustrative purposes and not for limitation, in order to provide a thorough understanding of the embodiments of the invention. However, those skilled in the art will understand that the invention can be implemented in other embodiments without these specific details. In other instances, detailed descriptions of well-known systems, apparatuses, circuits, and methods are omitted so as not to obscure the description of the invention with unnecessary detail.

[0021] This invention proposes a method for fabricating an implantable pressure sensor, the method comprising the following steps: Step S1: Prepare zwitterionic conductive hydrogels using zwitterionic monomers; Specifically, the preparation process in step S1 includes adding a crosslinking agent and an initiator to a polymerizable zwitterionic monomer, mixing them evenly, placing them in a mold, and polymerizing them under ultraviolet light to obtain a zwitterionic hydrogel; the polymerizable zwitterionic monomer includes 3-[[2-(methacryloyloxy)ethyl]dimethylammonium]propionate, [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide, and 2-methacryloyloxyethyl phosphocholine.

[0022] Specifically, the crosslinking agent is polyethylene glycol diacrylate (PEGDA), and the initiator is 2-hydroxy-2-methyl-1-[4-(2-hydroxyethoxy)phenyl]-1-propanone (Irgacure 2959). The degree of crosslinking of the zwitterionic hydrogel is controlled by adjusting the amount of crosslinking agent, thereby adjusting the mechanical properties and ion conductivity of the hydrogel.

[0023] Hydrogels with different degrees of crosslinking were prepared by adjusting the amount of crosslinking agent to optimize their pressure sensing sensitivity; preferably, the mass ratio of zwitterionic monomer, crosslinking agent and initiator was 2.1:0.025:0.01.

[0024] Step S2 involves copolymerizing acrylamide monomers and acrylate monomers as two components to prepare a microphase-separated copolymer hydrogel.

[0025] This invention prepares a prepolymer solution by adding a crosslinking agent, initiator and catalyst to acrylamide monomers and acrylate monomers; after sufficient polymerization to obtain a hydrogel, it is immersed in pure water to remove unreacted chemicals and reagents, thereby obtaining a microphase-separated dielectric gel.

[0026] Specifically, the acrylamide monomers include one or more combinations of acrylamide, methacrylamide, N,N-dimethylacrylamide, N-hydroxyethylacrylamide, N-isopropylacrylamide and their salts; the dielectric properties, compressibility and wet adhesion properties of the microphase separation copolymer hydrogel are controlled by adjusting the ratio of acrylamide monomers to acrylate monomers and the polymerization temperature.

[0027] Specifically, in step S2, the crosslinking agent is polyethylene glycol diacrylate (PEGDA), the initiator is ammonium persulfate (APS), and the catalyst is tetramethylethylenediamine (TEMED); preferably, the ratio of acrylamide monomer to acrylate monomer is 2:1, and the optimal reaction temperature is 90°C.

[0028] Step S3: The zwitterionic conductive hydrogel obtained in step S1 is used as the upper electrode layer and the lower electrode layer, and the microphase-separated copolymer hydrogel obtained in step S2 is used as the dielectric layer. The layers are assembled to obtain a three-layer structure device. Step S4: An encapsulation layer is provided on the outside of the three-layer structure device to obtain an implantable pressure sensor.

[0029] Specifically, the encapsulation layer is an electrospun aliphatic polyester material, and the material is bonded to the hydrogel layer through a mechanically interlocking interface.

[0030] This invention selects aliphatic polyester materials with excellent biocompatibility as electrospinning raw materials; the aliphatic polyester materials include, but are not limited to, polycaprolactone, polylactic acid, lactic acid-glycolic acid copolymer and polyglycolic acid.

[0031] By adjusting process parameters such as spinning solution concentration, applied voltage, ambient temperature and humidity, the porosity and mechanical properties of electrospun fiber membranes can be controlled and adjusted. Preferably, the spinning solution concentration is 20wt%, the applied voltage is 20kV, the ambient temperature is 25℃ and the relative humidity is 50%.

[0032] The present invention also proposes an implantable pressure sensor, the sensor comprising an electrode layer, a dielectric layer and an encapsulation layer; wherein the upper and lower electrode layers are both zwitterionic conductive hydrogels, the dielectric layer is a microphase-separated copolymer hydrogel, and the encapsulation layer is an aliphatic polyester material.

[0033] The present invention proposes the application of an implantable pressure sensor in physiological signal monitoring, wherein the physiological signals include tissue pressure signals, visceral motion signals, etc.

[0034] Example 1 The polymerizable zwitterionic monomer is carboxybenzene methacrylate (CBMA) monomer, and its preparation process is as follows: First, the reaction flask was purged with nitrogen to replace the air in the apparatus and maintain an inert atmosphere. Then, anhydrous acetone (300 mL) and N,N-dimethylaminoethyl methacrylate (60 mL) were added to the flask and stirred at room temperature until completely dissolved. The mixture was then stirred continuously at 0°C for 2 h. After that, β-propiolactone (21.82 mL, 0.35 mol) was slowly added dropwise to the flask using a constant pressure dropping funnel. The product was collected after reacting at room temperature for 2 h. The product was thoroughly washed with diethyl ether and then placed in a vacuum drying oven. After drying for 24 h, carboxybetaine methacrylate (CBMA) monomer was obtained.

[0035] The zwitterionic hydrogel polycarboxylated betaine methacrylate (PCBMA) was prepared using CBMA, and the specific preparation process is as follows: First, 2.1 g of CBMA was uniformly dissolved in 6 g of deionized water. Then, 0.025 g of crosslinking agent PEGDA and 0.01 g of initiator Irgacure 2959 were added sequentially, and the mixture was stirred for 5 min followed by ultrasonic degassing. The resulting solution was then transferred to a 1.0 mm thick mold using a dropper and irradiated under ultraviolet light for 30 min to obtain hydrogel PCBMA. Figure 1 As shown in the figure, the mass percentages of monomer CBMA, crosslinking agent PEGDA, and initiator Irgacure 2959 were 25.81%, 0.31%, and 0.12%, respectively (all rounded to two decimal places). The resulting hydrogel exhibited the best performance, so this ratio was used in subsequent experiments. Based on the same preparation process, hydrogels with different degrees of crosslinking were further prepared by adjusting the amount of crosslinking agent PEGDA added (Table 1).

[0036] Table 1. A series of PCBMA hydrogels with different degrees of crosslinking

[0037] Example 2 Microphase-separated copolymer hydrogels were prepared using the following method.

[0038] First, 2.1 g of hydroxyethyl methacrylate (HEMA) and acrylamide (AM) were dissolved in 7.2 g of propylene carbonate (PC) and stirred thoroughly until a homogeneous solution was formed. Then, 0.0125 g of crosslinking agent PEGDA, 0.01 g of initiator APS, and 0.01 g of catalyst TEMED were added to the solution in sequence, and the mixture was stirred for 5 min to ensure thorough mixing. The solution was then ultrasonically treated to remove air bubbles. Afterward, the resulting precursor solution was transferred to a mold with a thickness of 1.0 mm using a dropper and reacted in an oven for 6 h to obtain P(HEMA-co-AM) copolymer hydrogel. In this embodiment, the mass percentages of crosslinking agent PEGDA, initiator APS, and catalyst TEMED were fixed at 0.13%, 0.11%, and 0.11%, respectively (all rounded to two decimal places).

[0039] Using the same preparation process, a series of hydrogel samples with different compositions were prepared by adjusting the mass ratio of HEMA to AM (Table 2). Multiple experiments revealed that when the weight ratio of HEMA to AM was greater than 2:1, significant phase separation occurred in the system, resulting in a multiphase structure. Figure 2 As shown, considering the swelling properties of hydrogels, the optimal mass ratio of HEMA to AM was determined to be 2:1, and this ratio was used in all subsequent experiments.

[0040] Table 2. Copolymer hydrogels with different monomer ratios

[0041] During the preparation process, it was found that temperature has a significant impact on the adhesion properties of hydrogels. Therefore, in the preparation method of this invention, hydrogel samples under different conditions were prepared by controlling the reaction temperature to systematically study the effect of temperature on its adhesion properties and optimize the preparation parameters, thereby obtaining hydrogel materials with excellent adhesion properties (Table 3). Combined with adhesion force tests, the optimal reaction temperature of the above system was determined to be 90℃, so all subsequent experiments were conducted under this condition.

[0042] Table 3. Copolymer hydrogels under different temperature conditions

[0043] Example 3 The preparation process of polycaprolactone (PCL) electrospun film is as follows: Using PCL as the solute (20 wt%), and a mixed solution of acetone and N,N-dimethylformamide (DMF) in a volume ratio of 4:1, the total volume was 35 mL, comprising 28 mL acetone, 7 mL DMF, and 7 g PCL. During electrospinning, the applied voltage was 20 kV (18 kV at the positive electrode and 2 kV at the negative electrode), the solution feed rate was 1 mL / h, the ambient temperature was 25℃, the relative humidity was 50%, the spinning distance was 15 cm, and the needle used was 21G. Figure 3 As shown, electrospinning can produce uniform films with a certain porous structure, providing good mechanical support and biocompatibility.

[0044] Example 4 The fabrication process of the implantable pressure sensor is as follows: like Figure 4 As shown, the sensor has a three-layer parallel plate structure, including an upper electrode layer 1, a lower electrode layer 2, and a middle dielectric layer 3; an encapsulation layer 4 is provided on the outside of the structure; the fabrication process adopts a layer-by-layer construction method for assembly, and the wet adhesion characteristics of the dielectric layer hydrogel are used to achieve a tight bond between the layers. Finally, the overall structure is encapsulated by a PCL electrospun film.

[0045] (1) Electrode layer The upper and lower electrode plates are made of polycarboxybenzene methacrylate (PCBMA) hydrogel. This material has force-induced ion generation characteristics. Under the action of external force, water molecules in the system dissociate and generate free-migrating hydroxide ions, thereby causing a significant change in the dielectric constant of the material.

[0046] (2) Dielectric layer The intermediate layer 3 is a P(HEMA-co-AM) microphase separation copolymer hydrogel, which is prepared by copolymerizing HEMA and AM to form a stable microphase separation structure in the system. It has the characteristics of high dielectric constant, excellent compressive strength, and low mechanical hysteresis. At the same time, by controlling the preparation temperature, it can be endowed with excellent wet adhesion properties.

[0047] (3) Encapsulation layer 4: The encapsulation layer is made of polycaprolactone (PCL) film prepared by electrospinning process, and is combined with the hydrogel through mechanical interlocking interface structure, thereby improving the robustness and long-term stability of the encapsulation structure.

[0048] In this invention, the ionic conductivity of the zwitterionic hydrogel polycarboxybenzene methacrylate (PCBMA) was determined by AC impedance spectroscopy, and the specific steps are as follows: (1) Sample preparation The synthesized PCBMA hydrogel is cut into regular shapes (preferably circular pieces) and its initial size is measured; the sample is immersed in deionized water until it is fully swollen and balanced (usually not less than 24 h), and then the excess water on the surface is gently absorbed with filter paper for later use.

[0049] (2) Construction of the testing system Two test systems were constructed: (a) a pure water system, in which the hydrogel sample after swelling equilibrium was placed in deionized water; and (b) a salt solution system, in which the hydrogel sample was immersed in physiological saline (0.9% NaCl) until ion exchange equilibrium was reached.

[0050] (3) Conductivity measurement AC impedance testing was performed using an electrochemical workstation (such as CHI, Autolab, or equivalent equipment). During testing, the hydrogel sample was clamped between two conductive electrodes (preferably stainless steel or platinum electrodes) to ensure close contact between the electrodes and the sample. The sample thickness and the effective contact area of ​​the electrodes were recorded. A small-amplitude AC voltage (preferably 5–10 mV) was applied at room temperature, with a scanning frequency range of 1 Hz to 10 Hz. 6 Hz, to obtain the impedance spectrum (Nyquist plot); the resistance of the sample is obtained through the intersection of the high-frequency region and the real axis.

[0051] The ionic conductivity of hydrogels can be calculated using the following formula. :

[0052] In the formula, The sample thickness is in cm. Sample resistance, in Ω; S is the effective contact area of ​​the electrode, in cm².

[0053] The ionic conductivity of PCBMA hydrogel, measured using the above method, was 2.1 × 10⁻⁶ in both pure water and NaCl solution systems. -4 S / cm and 1.0×10 -2 The S / cm ratio indicates that this zwitterionic hydrogel has excellent ion conductivity.

[0054] The dielectric constant of the P(HEMA-co-AM) microphase-separated copolymer hydrogel was determined using an electrochemical workstation combined with AC impedance spectroscopy. The specific steps are as follows: (1) Sample preparation The prepared PCBMA hydrogel was cut into uniform sheet samples; before testing, the samples were equilibrated to stabilize their water content; the sample thickness was measured using a vernier caliper or thickness gauge.

[0055] (2) Electrode assembly Place the hydrogel sample to be tested between two parallel electrodes, preferably stainless steel, gold, or platinum electrodes; adjust the electrode assembly to ensure full contact between the sample and the upper and lower electrodes, and ensure the stability of the contact interface during the test; record the effective contact area of ​​the electrodes.

[0056] (3) Apply pressure and perform impedance testing: The assembled sample was subjected to AC impedance testing using an electrochemical workstation; the test frequency was set to 100 kHz, and the AC excitation voltage was 5–10 mV; the complex impedance parameters of the sample, including the real part of the impedance, were recorded at various set pressures. ) and the imaginary part of impedance ( Based on the impedance parameters obtained from the test, the dielectric constant of the hydrogel was calculated according to the following formula. The P(HEMA-co-AM) copolymer gel exhibits a high dielectric constant (≈38 at 100 kHz).

[0057] In the formula, ε is the dielectric constant, ω is the angular frequency in rad / s, and ε₀ is the vacuum dielectric constant. This represents the effective contact area of ​​the electrode, expressed in cm². The thickness of the sample is in cm.

[0058] In this invention, the mechanical properties of the hydrogel are characterized using a universal testing machine, and the compressive modulus is obtained through uniaxial compression testing. The specific steps are as follows: (1) Sample preparation The prepared hydrogel was cut into cylindrical samples with a diameter of 1 cm. Before testing, the samples were placed in a deionized water equilibrium environment to ensure that the water content inside the samples was consistent. The initial height h and cross-sectional area A of the samples were measured using vernier calipers.

[0059] (2) Instrument preparation The compressibility of the sample was tested using a universal testing machine with a parallel pressure plate fixture. The sample was placed between the upper and lower pressure plates and its position was adjusted so that the center of the sample coincided with the loading axis to ensure uniform force during compression.

[0060] (3) Compression test Under room temperature conditions, a uniaxial compression test is performed on the sample at a constant loading rate, preferably 5 mm / min. During the test, the load-displacement curve is continuously recorded until the sample reaches the set compressive strain to obtain a complete compressive mechanical response curve. The slope of the initial linear region of the stress-strain curve is taken as the compressive modulus of the sample. Preferably, a linear segment with a strain range of 5% to 15% is selected for fitting, and the resulting slope is the compressive modulus.

[0061] Calculations show that the elastic modulus of PCBMA hydrogel is approximately 45 kPa, which is close to the mechanical properties of human soft tissue, thus improving its compatibility with biological tissues. The elastic modulus of P(HEMA-co-AM) hydrogel is 90 kPa, indicating that it has high resistance to compressive deformation and structural stability under pressure.

[0062] In this invention, the interfacial adhesion properties between zwitterionic hydrogel PCBMA and P(HEMA-co-AM) hydrogel are characterized using the overlap shear method, with the specific steps as follows: (1) Sample preparation The prepared PCBMA hydrogel and P(HEMA-co-AM) hydrogel were cut into sheet samples of the same size, preferably with a length of 50 mm, a width of 10 mm, and a thickness of 2 mm. Before testing, the samples were equilibrated to ensure that the surface water content remained consistent.

[0063] (2) Interface construction The PCBMA hydrogel sample and the P(HEMA-co-AM) hydrogel sample are bonded together with a predetermined overlap area, preferably 10 mm × 10 mm. During the bonding process, a certain pressure can be applied at the interface and maintained for a certain time to ensure that the two hydrogels are in full contact and form a stable interface. It is preferred to apply slight pressure for 1 min.

[0064] (3) Instrument installation The adhesion force of the bonded composite sample was tested using a universal testing machine. The two ends of the PCBMA hydrogel and P(HEMA-co-AM) hydrogel were fixed on the upper and lower clamps of the testing machine, respectively, so that the tensile direction was parallel to the bonding interface to avoid sample displacement or uneven stress during the test. Shear separation test was performed at a constant tensile rate under room temperature conditions, with a preferred tensile rate of 5 mm / min. The load-displacement curve was continuously recorded during the test until the two hydrogels separated at the interface.

[0065] (4) Calculation of adhesion strength: Take the maximum separation load F recorded during the test. max As the interface failure load, the interfacial adhesion strength is calculated using the following formula. :

[0066] In the formula, τ is the interfacial adhesion strength, with units of kPa, and F max The maximum separation load is in N, and A is the interface contact area in cm².

[0067] The interfacial shear strength between hydrogel PCBMA and P(HEMA-co-AM) can reach 10 kPa, indicating that it has good interfacial bonding performance in flexible devices and layered composite structures; for example Figure 5 The P(HEMA-co-AM) hydrogel shown has excellent apparent adhesion properties and can be directly adhered to the surface of substrates such as paper, plastic, fiber, and metal.

[0068] In this invention, the degradation performance of zwitterionic hydrogel PCBMA and dielectric adhesive hydrogel P (HEMA-co-AM) was evaluated using the mass loss method, with the specific steps as follows: (1) Sample preparation The prepared PCBMA and P(HEMA-co-AM) hydrogels were cut into block samples of uniform size and weight. The samples were placed in deionized water and allowed to swell to equilibrium. The samples were then removed and the surface moisture was gently absorbed with filter paper. The initial mass of the samples was recorded.

[0069] (2) Construction of degradation system The weighed hydrogel sample was placed in a pre-prepared degradation medium, which was a phosphate buffer solution containing lipase; degradation was carried out at 37°C to simulate the physiological environment; to maintain enzyme activity and solution stability in the degradation system, the degradation medium was preferably replaced every 24 hours; the sample was taken out at the set time interval; the sample surface was gently rinsed with deionized water to remove residual substances, and then the surface moisture was absorbed with filter paper.

[0070] (3) Degradation rate determination The extracted sample was weighed, and its mass at each time point was recorded as follows: The degradation rate of the hydrogel is calculated using the following formula:

[0071] In the formula, The initial mass of the sample is expressed in grams. The mass of the sample at time t is expressed in grams.

[0072] Using the methods described above, Figure 6 The degradation behavior and mass loss of PCBMA and P(HEMA-co-AM) hydrogels in simulated physiological environments were shown, and the results indicate that both PCBMA and P(HEMA-co-AM) hydrogels have degradable properties.

[0073] In this invention, mouse fibroblasts (NIH / 3T3) were used to evaluate the cell compatibility of all components of the sensor. The specific steps are as follows: (1) Preparation of exudate All components of the sensor were immersed in DMEM medium for 24 hours, and the culture medium after immersion was collected as exudate for later use.

[0074] (2) Cell treatment NIH / 3T3 cells cultured for 36 h were digested with trypsin, then collected by centrifugation and resuspended in culture medium to prepare a homogeneous cell suspension.

[0075] (3) Cell inoculation and culture 200 μL of the cell suspension was seeded into a 48-well plate and incubated in an incubator.

[0076] (4) Group processing The culture medium was replaced with a medium containing the exudate as the experimental group; the control group used conventional DMEM medium without the exudate, and the cells in each group were incubated for another 24 hours.

[0077] (6) Cell morphology observation The morphology and growth status of cells in each group were observed using a fluorescence inverted microscope.

[0078] like Figure 7 As shown, there were no significant differences in cell morphology and proliferation between the experimental group and the control group. The absorbance test results also showed no significant difference in cell viability, indicating that the components of the sensor have good cell compatibility and no significant adverse effects on cells.

[0079] In this invention, the pressure response performance and stability testing of the hydrogel-based capacitive pressure sensor are specifically carried out through the following steps: (1) Connection of test device The upper and lower electrodes of the pressure sensor are connected to the positive and negative terminals of the LCR digital bridge, respectively, to construct a capacitive signal acquisition system. (2) Stress loading test like Figure 8 As shown, different external pressures are applied to the pressure sensor, and its capacitance change value is recorded in real time during the pressure loading process, and the corresponding relationship data between pressure and capacitance response is obtained.

[0080] (3) Sensitivity curve test Within a preset pressure range, pressure is applied in stages, and the corresponding capacitance changes are recorded. Based on the obtained data, the sensor's sensitivity is calculated, and the relationship between applied pressure and capacitance changes is analyzed to evaluate its linear response characteristics. Figure 9As shown; the test results indicate that the pressure sensor possesses excellent pressure sensitivity characteristics, and the relative capacitance change rate (ΔC / C0) exhibits a piecewise linear relationship with the applied pressure: the sensitivity S1 in the low-pressure range (0–1 kPa) is 0.0752 kPa. - ¹, The sensitivity S² in the high-pressure range (1–4 kPa) is 0.0348 kPa. - ¹.

[0081] (4) Cyclic stability test Under cyclic loading conditions, external pressure is repeatedly applied and released to the pressure sensor, and the changes in its output signal are continuously recorded to evaluate its signal stability during multiple loading processes. Figure 10 As shown.

[0082] Test results show that the pressure sensor exhibits high sensitivity within the pressure range of 0.08–1 kPa, with a sensitivity reaching 0.075 kPa. - ¹, and the relationship between pressure and capacitance changes is approximately linear; at the same time, under cyclic loading and long-term loading conditions, its output signal does not show obvious fluctuations or drift; after the pressure is released, the capacitance value can recover to the initial state, indicating that the sensor has good stability, recoverability and repeatability.

[0083] In this invention, the sensor is implanted into target tissue sites (such as the gut microbiota, heart, bladder, blood vessels, soft tissue, etc.) within experimental animals to collect and analyze signals generated during tissue pressure changes in real time. The specific steps are as follows: (1) Preparation of experimental animals Healthy laboratory animals were selected, acclimatized before the experiment, and preoperative treatment was performed as needed to reduce the impact of the internal environment on the surgical and monitoring results.

[0084] (2) Anesthesia and surgical site management The experimental animal was anesthetized, and after complete anesthesia, it was fixed on the operating table. The target implantation area was prepared and disinfected. A small incision was made along the predetermined position, and the skin and muscle layers were separated layer by layer to expose the target tissue.

[0085] (3) Sensor implantation and fixation The sensor is fixed to the surface or interior of the target tissue using medical sutures or bio-adhesive to prevent significant displacement or detachment during tissue movement. A lead wire is led out subcutaneously and connected to the signal acquisition device.

[0086] (4) Acquisition of internal pressure signals Once the experimental animals have stabilized, the signal acquisition system is activated to continuously record pressure changes in the tissues under natural conditions or external stimuli, acquiring the corresponding real-time response signals. Depending on experimental requirements, tissue pressure changes can be induced by applying mechanical stimulation, drug administration, or other physiological modulation methods, while simultaneously recording the amplitude, frequency, and periodic changes of the sensor output signals before and after stimulation.

[0087] like Figure 11 As shown, after the sensor is implanted in the body, it can record the dynamic signals generated by changes in tissue pressure in real time. Under different loading frequencies (2, 3 and 4 cpm), the collected signals have obvious periodic changes, and the changes in amplitude and frequency can reflect the changes in tissue pressure state, indicating that the sensor has good in vivo pressure monitoring capabilities and can be used to monitor physiological rhythms (such as respiration or gastrointestinal peristalsis).

[0088] In the above embodiments, the descriptions of each embodiment have different focuses. For parts that are not described in detail or recorded in a certain embodiment, please refer to the relevant descriptions of other embodiments.

[0089] The above-described embodiments are only used to illustrate the technical solutions of the present invention, and are not intended to limit it. Although the present invention has been described in detail with reference to the foregoing embodiments, those skilled in the art should understand that modifications can still be made to the technical solutions described in the foregoing embodiments, or equivalent substitutions can be made to some of the technical features. Such modifications or substitutions do not cause the essence of the corresponding technical solutions to deviate from the spirit and scope of the technical solutions of the embodiments of the present invention, and should all be included within the protection scope of the present invention.

Claims

1. A method for fabricating an implantable pressure sensor, characterized in that, The preparation method includes the following steps: Step S1: Prepare zwitterionic conductive hydrogels using zwitterionic monomers; Step S2: A microphase-separated copolymer hydrogel is prepared by copolymerizing acrylamide monomers and acrylate monomers as two components. Step S3: The zwitterionic conductive hydrogel obtained in step S1 is used as the upper electrode layer and the lower electrode layer, and the microphase-separated copolymer hydrogel obtained in step S2 is used as the dielectric layer. The layers are assembled to obtain a three-layer structure device. Step S4: An encapsulation layer is provided on the outside of the three-layer structure device to obtain an implantable pressure sensor.

2. The preparation method according to claim 1, wherein in step S2, the acrylamide monomer includes one or more combinations of acrylamide, methacrylamide, N,N-dimethylacrylamide, N-hydroxyethylacrylamide, N-isopropylacrylamide and their salts.

3. According to the preparation method of claim 1, in step S4, the encapsulation layer is an electrospun aliphatic polyester material.

4. The preparation method according to claim 3, wherein the aliphatic polyester material includes, but is not limited to, polycaprolactone, polylactic acid, lactic acid-glycolic acid copolymer, and polyglycolic acid.

5. The preparation method according to claim 1, wherein in step S1, the zwitterionic monomers include 3-[[2-(methacryloyloxy)ethyl]dimethylammonium]propionate, [2-(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl)ammonium hydroxide and 2-methacryloyloxyethyl phosphocholine.

6. According to the preparation method of claim 1, in step S1, a prepolymer solution is prepared by adding a crosslinking agent and an initiator to the zwitterionic monomer; after sufficient polymerization to obtain a hydrogel, it is soaked in pure water to remove unreacted chemical substances and reagents, thereby obtaining a zwitterionic conductive gel.

7. According to the preparation method of claim 1, in step S2, when copolymerizing acrylamide monomers and acrylate monomers as two components, a crosslinking agent, initiator and catalyst are added to prepare a prepolymer solution; after sufficient polymerization to obtain a hydrogel, it is soaked in pure water to remove unreacted chemical substances and reagents, and a microphase-separated copolymer hydrogel is obtained.

8. The preparation method according to claim 6, wherein the initiator in step S1 is 2-hydroxy-2-methyl-1-[4-(2-hydroxyethoxy)phenyl]-1-propanone.

9. The preparation method according to claim 7, wherein the initiator in step S2 is ammonium persulfate.

10. The preparation method according to claim 6, wherein the crosslinking agent is polyethylene glycol diacrylate.

11. The preparation method according to claim 7, wherein the catalyst is tetramethylethylenediamine.

12. According to the preparation method of claim 7, in step S2, the mass ratio of acrylamide monomer to acrylate monomer is 2:

1.

13. An implantable pressure sensor prepared using the preparation method according to any one of claims 1-12, characterized in that, The sensor includes an upper electrode layer (1), a lower electrode layer (2), a dielectric layer (3), and an encapsulation layer (4); wherein the upper electrode layer (1) and the lower electrode layer (2) are both zwitterionic conductive hydrogels, the dielectric layer (3) is a microphase-separated copolymer hydrogel, and the encapsulation layer (4) is an aliphatic polyester material.

14. The application of the implantable pressure sensor according to claim 13 in physiological signal monitoring, characterized in that, The physiological signals include tissue pressure signals and visceral movement signals.

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