A nanometer-pore transistor biomolecule sensor based on surface acoustic wave confinement and a preparation method thereof

Abstract

The application discloses a kind of nanometer hole transistor biomolecule sensors based on surface acoustic wave constraint and preparation method thereof.The sensor includes substrate, silicon nitride film layer, detection layer which are sequentially stacked, first passivation layer covering detection layer, surface acoustic wave excitation structure arranged on first passivation layer, second passivation layer covering surface acoustic wave excitation structure, nanometer hole penetrating through the sensor, and hydrophilic gate dielectric oxide layer covering the inner circumferential surface of nanometer hole.The surface wave excitation structure is used to excite surface acoustic wave which is focused in nanometer hole and produces blocking effect on biomolecule, including piezoelectric film, and interdigital electrode layer arranged on piezoelectric film.The application realizes broad-spectrum detection of multiple biomolecules without modifying specific biological capture layer through acoustic streaming effect of surface acoustic wave, charge induction of transistor and confinement effect of nanometer hole.

Description

Technical Field

[0001] This invention belongs to the field of nanomaterials and biosensing technology, specifically relating to a nanoporous transistor biomolecular sensor based on surface acoustic wave confinement and its preparation method. Background Technology

[0002] In the field of biomedical detection and analysis, biosensors based on biological or solid-state nanopores have shown great application potential due to their superior performance in detection and analysis at the single biomolecule level, especially in key areas such as gene sequencing, biochemical analysis, and early disease detection.

[0003] Currently, combining precise control of single molecules through nanopores with field-effect charge sensing technology to achieve charge sensing at the single biomolecule level has become a cutting-edge research direction in solid-state nanopore devices. The main sensing principle of these devices is based on the reduction in ionic conductivity caused by the blockage of nanopores when molecules in solution pass through them. To further improve device performance and application range, introducing more sensing methods at the front end of the nanopore, enabling the simultaneous detection of other physical properties of biomolecules, would undoubtedly greatly enhance the device's recognition capabilities and open up a wider range of applications.

[0004] With the continuous development of microelectronics and microelectromechanical systems (MEMS) technology, it has become possible to integrate electronically functional micro / nano structures into solid-state nanopores. However, it should be noted that most existing technologies rely on the passive diffusion of biomolecules through the pores, resulting in excessively fast perforation speeds and short signal capture time windows, which severely restricts the accuracy and signal-to-noise ratio of detection. Summary of the Invention

[0005] This invention provides a nanopore transistor biomolecular sensor based on surface acoustic wave confinement and its preparation method. This sensor achieves precise control and high-sensitivity detection at the single-molecule level by combining the acoustic flow effect of surface acoustic waves, the charge sensing of transistors, and the confinement effect of nanopores. It avoids dependence on biological probes and is suitable for high-sensitivity biomolecular analysis such as deoxyribonucleic acid (DNA) sequencing, protein detection, and virus screening.

[0006] In a first aspect, the present invention provides a nanopore transistor biomolecular sensor based on surface acoustic wave (SAW) confinement, comprising a substrate, a silicon nitride thin film layer, a detection layer, and a nanopore penetrating the sensor, which are sequentially stacked. The nanopore transistor biomolecular sensor further includes a first passivation layer covering the detection layer, a SAW excitation structure disposed on the first passivation layer, a second passivation layer covering the SAW excitation structure, and a hydrophilic gate dielectric oxide layer covering the inner peripheral surface of the nanopore.

[0007] The detection layer includes a source electrode, a drain electrode, and a carbon nanotube channel structure connecting the source electrode and the drain electrode. The nanopores pass through the carbon nanotube channel structure. The surface wave excitation structure, used to excite surface acoustic waves focused in the nanopores to produce a blocking effect on biomolecules, includes a piezoelectric thin film and an interdigitated electrode layer disposed on the piezoelectric thin film.

[0008] This invention introduces surface acoustic wave (SAW) technology at the front end of a nanopore sensing structure. SAW excited by interdigitated electrodes can generate a powerful acoustic flow in a microfluidic environment, thereby actively driving and precisely controlling the movement of biomolecules (such as DNA) into the nanopore. This active manipulation method effectively slows down the molecular transit velocity and significantly prolongs their residence time within the nanopore, providing a sufficient time window for acquiring high signal-to-noise ratio electrical signals.

[0009] Preferably, the carbon nanotube channel structure uses single-walled semiconductor carbon nanotubes; the side of the carbon nanotube channel structure facing away from the silicon nitride thin film layer has an exposed window surrounding the nanopore, so that the nanopore forms a stepped pore structure.

[0010] Preferably, the interdigitated electrode layer includes two annular interdigitated electrode units symmetrically arranged on both sides of the nanopore.

[0011] Preferably, the hydrophilic gate dielectric oxide layer is made of aluminum oxide. The hydrophilic gate dielectric oxide layer 8 completely covers the cross-sections of the first passivation layer and the second passivation layer in the nanopore, and also covers the outer surface of the second passivation layer.

[0012] Preferably, the substrate has a through-hole structure, exposing silicon nitride windows at the corresponding nanopore positions on the silicon nitride thin film layer.

[0013] Preferably, the diameter of the nanopore aligned with the detection layer 4 is 5 nm to 10 nm.

[0014] Secondly, the present invention provides a biomolecule detection method that utilizes the aforementioned nanoporous transistor biomolecule sensor. The biomolecule detection method includes:

[0015] An electrolyte containing the target biomolecules is introduced into a nanopore. A gate voltage is applied to the electrolyte, and source-drain voltages are applied to the source and drain electrodes. An excitation signal is applied to the interdigitated electrode layer, causing the piezoelectric film to excite surface acoustic waves, which impede the biomolecules passing through the nanopore. The target biomolecules are detected based on the changes in the source and drain electrode currents.

[0016] Preferably, the input power of the surface acoustic wave is 26dBm to 28dBm.

[0017] Thirdly, the present invention provides a sensor manufacturing method, comprising:

[0018] Silicon nitride thin films are grown on both sides of a silicon wafer as a substrate.

[0019] Marker points are patterned on the front side of the resulting substrate, and silicon nitride is removed by patterning on the back side to expose the etched window area.

[0020] The substrate on the back side of the substrate is etched. The etching depth is 80% to 90% of the substrate thickness.

[0021] The substrate was immersed in a toluene dispersion of carbon nanotubes, allowing the carbon nanotubes to deposit onto the substrate surface. After heating and curing the carbon nanotubes, excess carbon nanotubes were etched away to form a carbon nanotube channel structure.

[0022] After forming the source electrode and drain electrode, which connect the carbon nanotube channel structure, on the front side of the substrate, a first passivation layer is applied.

[0023] Zinc oxide is grown by magnetron sputtering on the first passivation layer to form a piezoelectric thin film.

[0024] After forming an interdigitated electrode layer on the piezoelectric thin film, a second passivation layer is applied.

[0025] The first passivation layer, the piezoelectric film, and the second passivation layer are etched away in the area surrounding the nanopore.

[0026] Deep silicon etching was performed on the back side of the substrate to expose a silicon nitride window.

[0027] After creating the nanopores, aluminum oxide is grown on the inner surface of the nanopores as a hydrophilic gate medium oxide layer 8.

[0028] Fourthly, the present invention provides a biomolecule detection device, comprising a circuit board 10, a sealing ring, clamps, and the aforementioned nanoporous transistor biomolecule sensor. The nanoporous transistor biomolecule sensor is fixed in the detection area of ​​the circuit board. The detection area of ​​the circuit board is clamped between two of the clamps.

[0029] Preferably, sealing rings are provided between the two clamps and both sides of the circuit board. The sealing rings surround the detection area of ​​the circuit board 10. The two clamps are fixed together and the sealing rings are pressurized.

[0030] Preferably, the circuit board also includes an interface area. A detection signal interface and an ultrasonic excitation signal interface are fixed to the interface area of ​​the circuit board. The source and drain electrodes of the nanoporous transistor biomolecular sensor are led out to the detection signal interface through a conductive layer on the circuit board. The interdigitated electrode layer of the nanoporous transistor biomolecular sensor is led out to the ultrasonic excitation signal interface through a conductive layer on the circuit board. The electrical connection structure between the conductive layer of the circuit board and the nanoporous transistor biomolecular sensor, as well as the conductive layer within the detection area, are all sealed and hydrophilicized.

[0031] The beneficial effects of this invention are as follows:

[0032] 1. This invention utilizes the acoustic flow effect of surface acoustic waves, the charge sensing of transistors, and the confinement effect of nanopores to achieve broad-spectrum detection of various biomolecules (such as DNA, proteins, and viruses) without the need for modification of specific biocapture layers. This avoids the complexity and selectivity limitations of coating preparation and significantly improves the versatility and practicality of the sensor.

[0033] 2. This invention generates a blocking effect on biomolecules by exciting and focusing surface acoustic waves on nanopores, thereby regulating the speed at which biomolecules pass through the nanopores. Furthermore, by adjusting the power and frequency of the surface acoustic waves, the acoustic flow intensity can be controlled to obtain the electrical signals of biomolecules at different passage speeds, providing unprecedented multidimensional information for studying molecular conformational changes and interaction dynamics.

[0034] 3. This invention utilizes the nano-vortices generated by surface acoustic waves under the confinement effect of nanopores to capture molecules at the front end of nanopores, promoting the unfolding of folded biomolecules and preventing nanopore blockage. It also helps to solve the problem of difficult signal interpretation caused by random molecular folding and crossing in traditional nanopore technology, greatly improving the consistency and interpretability of electrical signals, and laying the foundation for precise sequencing or structural analysis of biomolecules.

[0035] 4. The transistor-embedded nanopore sensor provided by this invention can improve the signal-to-noise ratio (SNR) of the current signal to more than 10 times that of traditional nanopores, significantly improving the detection sensitivity and accuracy of biomolecular perforation events. Combined with the nanomanipulation effect of surface acoustic waves, this invention can capture biomolecular perforation events in microseconds, providing strong technical support for real-time dynamic monitoring of biomolecular interactions, gene sequencing at the single-molecule level, and research on protein conformational changes.

[0036] 5. This invention, by coating the inner surface of the nanopore with a hydrophilic gate dielectric oxide layer, effectively prevents non-target molecules from adsorbing onto the sensitive area of ​​the transistor channel surface through electrostatic or hydrophobic interactions, allowing only target molecules actively driven to pass through the channel to generate signals, thus further improving the specificity and accuracy of detection. Simultaneously, the hydrophilic gate dielectric oxide layer helps to further reduce the nanopore size and enhance the confinement effect. Attached Figure Description

[0037] To more clearly illustrate the technical solutions in the embodiments of the present invention or the prior art, 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.

[0038] Figure 1 This is a schematic diagram of the structure of the nanopore transistor biomolecular sensor provided in Embodiment 1 of the present invention.

[0039] Figure 2 This is a schematic diagram of the detection layer and the interdigitated electrode layer in Embodiment 1 of the present invention.

[0040] Figure 3 This is a schematic diagram of the detection layer in Embodiment 1 of the present invention.

[0041] Figure 4 This is a schematic diagram of the structure of the interdigitated electrode layer in Embodiment 1 of the present invention.

[0042] Figure 5 This is a schematic flowchart of the preparation method provided in Embodiment 2 of the present invention.

[0043] Figure 6 This is an exploded schematic diagram of the detection device provided in Embodiment 3 of the present invention.

[0044] Figure 7 This is a graph showing the perforation response curves of DNA with different conformations in the detection method of Example 4 of the present invention.

[0045] Reference numerals: 1. Substrate; 2. Silicon nitride thin film layer; 2-1. Silicon nitride window; 3. First passivation layer; 4. Detection layer; 4-1. Source electrode; 4-2. Drain electrode; 4-3. Carbon nanotube channel structure; 5. Piezoelectric thin film; 6. Interdigitated electrode layer; 7. Second passivation layer; 8. Hydrophilic gate dielectric oxide layer; 9. Nanopore; 10. Circuit board; 10-1. Detection signal interface; 10-2. Ultrasonic excitation signal interface; 11. Sealing ring; 12. Fixture. Detailed Implementation

[0046] The present invention will be further described in detail below with reference to the embodiments. The following embodiments are explanations of the present invention, but the present invention is not limited to the following embodiments.

[0047] Example 1

[0048] like Figure 1 and Figure 2 As shown, a nanopore transistor biomolecular sensor based on surface acoustic wave confinement includes a substrate 1, a silicon nitride thin film layer 2, a first passivation layer 3, a detection layer 4, a piezoelectric thin film 5, an interdigitated electrode layer 6, a second passivation layer 7, and a hydrophilic gate dielectric oxide layer 8.

[0049] The substrate 1, silicon nitride thin film layer 2, first passivation layer 3, piezoelectric thin film 5, and second passivation layer 7 are sequentially stacked. A through-hole structure is formed on the substrate 1. This through-hole structure exposes the side surface of the silicon nitride thin film layer 2, forming a silicon nitride window 2-1. A nanopore 9, penetrating the silicon nitride thin film layer 2, first passivation layer 3, piezoelectric thin film 5, and second passivation layer 7, is formed in the center of the biomolecular sensor. The nanopore 9 is located at the center of the silicon nitride window 2-1. The axis of the nanopore 9 is perpendicular to the silicon nitride thin film layer 2. The piezoelectric thin film 5 and the interdigitated electrode layer 6 form a surface acoustic wave excitation structure.

[0050] The hydrophilic gate dielectric oxide layer 8 covers the outer side of the second passivation layer 7 (i.e., the side facing away from the second passivation layer 7) and the inner peripheral surface of the nanopore 9. The hydrophilic gate dielectric oxide layer 8 is used to isolate the carbon nanotube channel structure 4-3 in the detection layer 4 from the external electrolyte, and to encapsulate the hydrophobic passivation layer cross section in the nanopore 9, so that the inner surface of the nanopore 9 has higher hydrophilicity.

[0051] The detection layer 4 is embedded in the first passivation layer 3. The interdigitated electrode layer 6 is embedded in the second passivation layer 7. The first passivation layer 3 separates the detection layer 4 from the piezoelectric film 5. The second passivation layer 7 separates the interdigitated electrode layer 6 from the hydrophilic gate dielectric oxide layer 8, eliminating or suppressing the generation of induced electrical signals between the interdigitated electrode layer 6 and the external electrolyte.

[0052] like Figure 2 and Figure 3As shown, the detection layer 4 performs biomolecule detection using the field-effect transistor principle, and includes a source electrode 4-1, a drain electrode 4-2, and a carbon nanotube channel structure 4-3. The source electrode 4-1 and the drain electrode 4-2 are symmetrically arranged on both sides of the nanopore 9 and connected through the carbon nanotube channel structure 4-3. The nanopore 9 passes through the carbon nanotube channel structure 4-3. An exposure window is formed around the nanopore 9 on the side of the carbon nanotube channel structure 4-3 away from the silicon nitride thin film layer 2, and the exposed surface on the carbon nanotube channel structure 4-3 forms a gate sensing surface. This exposure window allows the nanopore (9) to form a stepped hole structure.

[0053] During the detection process, the electrolyte containing biomolecules acts as the gate medium and contacts the gate sensing surface. Under the gate voltage, the channel conductivity is modulated, thereby generating a detectable current disturbance between the source electrode 4-1 and the drain electrode 4-2.

[0054] like Figure 2 and Figure 4 As shown, the interdigitated electrode layer 6 includes two annular interdigitated electrode units. The annular interdigitated electrode units adopt a focused interdigitated electrode structure and are symmetrically arranged on both sides of the nanopore 9.

[0055] In some embodiments, the carbon nanotube channel structure 4-3 uses single-walled semiconductor carbon nanotubes with a purity greater than or equal to 99%.

[0056] In some embodiments, both the source electrode 4-1 and the drain electrode 4-2 employ a titanium, palladium, and gold three-layer structure, forming good contact with the carbon nanotube channel structure 4-3. In some further embodiments, the thicknesses of the titanium, palladium, and gold three-layer structure are 0.6 nm, 20 nm, and 30 nm, respectively; wherein titanium serves as the adhesive layer, and the high work function palladium and gold form ohmic contacts with the carbon nanotubes.

[0057] In some embodiments, the first passivation layer 3 and the second passivation layer 7 are both made of AR-P 5350 photoresist.

[0058] The nanoporous transistor biomolecular sensor provided in this embodiment detects biomolecules based on the principle of biomolecular charge sensing. The specific detection process is as follows:

[0059] During the detection process, the biomolecular sensor is immersed in an electrolyte containing biomolecules. A gate voltage is applied to the electrolyte, while a source-drain voltage is applied to the biomolecular sensor, keeping the detection layer 4 within the linear operating region of the transistor. A surface acoustic wave (SAW) excitation structure composed of a piezoelectric film 5 and an interdigitated electrode layer 6 is used to excite SAW waves with the nanopore as the focusing region. Biomolecules pass through the nanopore 9 under the synergistic effects of electroosmosis, electrophoresis, and sonication. The biomolecules carry charges in the electrolyte at a specific pH value, causing charge perturbation around the nanopore 9, thereby altering the source-drain current of the transistor. This allows for the simultaneous detection of ion current blocking signals and biomolecular charge perturbation signals. Specifically, the electrophoretic force pushes the biomolecules downwards through the pore, while the acoustic waves delay their passage upwards.

[0060] The process of exciting surface acoustic waves (SAWs) is as follows: A specific frequency excitation signal is applied to the interdigitated electrodes of the interdigitated electrode layer 6. Under the action of an alternating electric field, the piezoelectric material beneath the interdigitated electrodes undergoes periodic mechanical deformation. The periodic arrangement of the interdigitated electrodes causes the minute deformations excited beneath each interdigitated electrode to superimpose and interfere with each other on the surface of the piezoelectric film 5, ultimately forming a mechanical wave, i.e., a SAW, whose energy is concentrated and propagates on the surface of the piezoelectric film 5. The design of the focusing interdigitated electrodes concentrates the energy of the SAWs, confining the energy to a nanoscale focusing region.

[0061] The surface acoustic waves (SAWs) excited by the interdigitated electrodes propagate along the chip surface and generate a strong acoustic flow effect through coupling with the liquid. At the front end of the nanopore 9, the SAWs are converted into leakage Rayleigh wave mode when propagating from the solid to the solid-liquid interface. The energy of the leakage Rayleigh wave radiates into the liquid along a direction with a Rayleigh angle to the vertical direction, and the resulting upward component slows down the passage of biomolecules through the pore. The active control of the molecular passage process can be achieved by adjusting the intensity of the SAWs. At the entrance of the nanopore 9 (which is also the gate sensing surface), the SAWs generate nano-eddies in the confined environment of the nanopore 9. The shear force generated by the eddies promotes the opening of the folded state of biomolecules, allowing biomolecules to pass through the nanopore 9 in a near-linear state. The entrance of the nanopore 9 is the end of the nanopore 9 away from the substrate 1; the exit of the nanopore 9 is the end of the nanopore 9 closer to the substrate 1.

[0062] In some embodiments, the thickness of the carbon nanotube channel structure 4-3 is 1.3-10 nm.

[0063] In some embodiments, the length of the carbon nanotube channel structure 4-3 is 1-5 μm and the width is 1-10 μm.

[0064] In some embodiments, the contact width between the carbon nanotube channel structure 4-3 and the source electrode 4-1 and the drain electrode 4-2 is 3-5 μm, thereby achieving good ohmic contact between the source electrode 4-1, the drain electrode 4-2 and the carbon nanotube channel structure 4-3.

[0065] In some embodiments, the diameter of the nanopore 9 at the alignment position with the detection layer 4 is 5-10 nm.

[0066] In some embodiments, the silicon nitride window 2-1 is rectangular with a side length less than or equal to 100 μm. An excessively large silicon nitride window 2-1 will introduce significant noise that affects the detection of biomolecules, while an excessively small window will be difficult to find when fabricating nanopores 9.

[0067] In some embodiments, the hydrophilic gate dielectric oxide layer 8 is formed by magnetron sputtering or thermal oxidation.

[0068] In some embodiments, the thickness of the hydrophilic gate dielectric oxide layer 8 is 5-10 nm, which can reduce the diameter of the nanopore 9 while isolating the electrolyte solution from the transistor. In addition, the thinner hydrophilic gate dielectric oxide layer 8 can reduce the threshold voltage of the transistor.

[0069] In some embodiments, the piezoelectric film 5 is a zinc oxide piezoelectric film 5.

[0070] In some embodiments, the hydrophilic gate dielectric oxide layer 8 is made of aluminum oxide.

[0071] In some embodiments, the silicon nitride thin film layer 2 is a low-stress silicon nitride (Si3N4) thin film with a thickness of 10-20nm, which has both mechanical strength and insulation properties.

[0072] In some embodiments, the thickness of the piezoelectric film 5 is 2-5 μm.

[0073] In some embodiments, the frequency range of surface acoustic waves excited by the interdigitated electrode layer and piezoelectric thin film is 100MHz-2GHz.

[0074] In some embodiments, the thickness of the hydrophilic gate dielectric oxide layer is 5-25 nm.

[0075] Example 2

[0076] like Figure 5 As shown, a method for fabricating a nanoporous transistor biomolecular sensor based on surface acoustic wave confinement includes the following steps:

[0077] Step 1: Using (100) oriented, P-type doped low-resistivity silicon as substrate 1, a dense silicon nitride thin film layer 2 with a thickness of 10-20 nm is grown on both sides of the silicon by plasma-enhanced chemical vapor deposition (PECVD). The thickness of the substrate 1 is 300 μm.

[0078] Step 2: The front side of the substrate obtained in Step 1 is patterned using a lift-off process to form marker points (i.e., the MARK layer). 20nm titanium and 60nm gold are deposited by electron beam evaporation to ensure subsequent overlay and positioning.

[0079] Step 3: Pattern the etching window area on the back of the substrate. The reactive ion etching creates a square window with a side length of 470μm. When arraying the sensor chips on the wafer, the center distance between them is 1cm.

[0080] Step 4: Immerse the substrate obtained in Step 3 in a KOH solution with a concentration of 45% and a temperature of 85℃. The etching rate of silicon by the KOH solution is 0.5~0.6μm / min. Silicon nitride can remain intact and stable for a long time. The etching time is 8h, and the etching depth is 80%-90% of the initial thickness of substrate 1. 10%-20% of the silicon thickness is retained at the window position of substrate 1 as support in subsequent processing.

[0081] Step 5: Preparation of a short, high-purity single-walled carbon nanotube toluene dispersion. The preparation process is as follows: Add 10 mg of short, high-purity single-walled carbon nanotube powder and 10 mg of polymer PCz (poly[9-(1-octylnonyl)-9H-carbazole]) to 100 ml of toluene solution, and stir until the powder is completely submerged. Place the solution in an ultrasonic homogenizer in an ice bath environment for ultrasonication. The ultrasonic power is 300 W, and the switching time in one cycle is 3 s. Maintain the ice bath for a total of 30 min of ultrasonication to obtain a uniformly dispersed short, high-purity single-walled carbon nanotube toluene dispersion. The short, high-purity single-walled carbon nanotube powder is prepared by a floating catalytic method.

[0082] The carbon nanotube toluene dispersion was diluted to 10 μg / ml for later use. The substrate obtained in step four was immersed in the 10 μg / ml carbon nanotube toluene dispersion, and the carbon nanotubes were naturally deposited onto the substrate surface. The crystallization dish was sealed with tin foil to prevent toluene evaporation. After 10-24 hours, the substrate was removed and placed in a pure toluene solution for 10 minutes to remove loose carbon nanotubes, and gently shaken. The substrate surface was rinsed with toluene and isopropanol, dried with nitrogen, and baked on a hot plate at 150°C for 30 minutes to make the carbon nanotubes adhere more firmly. Photolithography and reactive ion etching (RIE) were used to remove excess carbon nanotubes to form carbon nanotube channel structures 4-3. The reactive ion etching used oxygen as the working gas, the RIE power was 60W, and the etching time was 30s. The deposited substrate was removed and annealed at 600°C for 30 minutes in an inert environment to remove the polymer.

[0083] Step 6: The front side of the substrate obtained in Step 5 is patterned using a lift-off process, and 0.6nm titanium, 20nm palladium, and 30nm gold are deposited by electron beam evaporation to form ohmic contacts with the carbon nanotube channels, forming source electrode 4-1 and drain electrode 4-2.

[0084] Step 7: Front-side photolithography patterning. A 1μm thick AR-P 5350 photoresist layer is applied to the front side of the substrate obtained in Step 6 as the first passivation layer 3. Hardening is performed at 150℃ for 10 minutes to isolate the source electrode 4-1, drain electrode 4-2 and the upper structure of the first passivation layer 3.

[0085] Step 8: A dense zinc oxide film of 3-5 μm is grown on the front side of the substrate obtained in step 7 by magnetron sputtering to form a piezoelectric thin film 5.

[0086] Step 9: The front side of the substrate obtained in Step 8 is patterned using a lift-off process, and 10nm titanium and 30nm gold are deposited by electron beam evaporation to form the interdigitated electrode layer 6.

[0087] Step 10: Front-side photolithography patterning. A 1μm thick AR-P 5350 photoresist layer is applied to the front side of the substrate obtained in step 9 as the second passivation layer 7. Harden the film at 150℃ for 10 minutes to isolate the interdigitated electrode layer 6 from the external electrolyte.

[0088] Step 11: On the front side of the substrate obtained in Step 10, 5% dilute hydrochloric acid is used for etching to remove the area around the first passivation layer 3, piezoelectric film 5, and second passivation layer 7 around the nanopore 9, exposing the pads of the carbon nanotube channel structure 4-3, source electrode 4-1, drain electrode 4-2, and interdigitated electrode layer 6. Deep silicon etching is then performed on the back side of the substrate using inductively coupled plasma etching (ICP) to remove the remaining silicon in the window, exposing a square silicon nitride film window with a side length of 100 μm. At this point, a carbon nanotube channel structure-silicon nitride bilayer structure is formed in the center of the substrate. Nanopores 9 with a diameter of 15-30 nm are then created through the carbon nanotube channel structure-silicon nitride bilayer structure using focused ion beam (FIB) technology.

[0089] Step 12: A 10-20 nm alumina layer was grown by magnetron sputtering as the growth hydrophilic gate medium oxide layer 8, which improved the hydrophilicity of the chip surface and reduced the size of the nanopores. The final nanopore diameter range was 5-10 nm.

[0090] Example 3

[0091] like Figure 6 As shown, a biomolecule detection device includes a circuit board 10, a sealing ring 11, a clamp 12, and a nanoporous transistor biomolecule sensor provided in Example 1.

[0092] The circuit board 10 has a detection area and an interface area. The detection area is circular. The interface area of ​​the circuit board 10 is fixed with a detection signal interface 10-1 and an ultrasonic excitation signal interface 10-2. The circuit board 10 is a printed circuit board.

[0093] The nanoporous transistor biomolecular sensor is fixed in the detection area of ​​the circuit board 10. The source electrode 4-1 and drain electrode 4-2 of the nanoporous transistor biomolecular sensor are led out to the detection signal interface 10-1 through the conductive layer on the circuit board 10. The interdigitated electrode layer 6 of the nanoporous transistor biomolecular sensor is led out to the ultrasonic excitation signal interface 10-2 through the conductive layer on the circuit board 10.

[0094] The conductive layer of the circuit substrate 10 and the electrical connection structure between the nanoporous transistor biomolecular sensor, as well as the conductive layer within the detection area, are all sealed and hydrophilicized.

[0095] There are two clamps 12. The detection area of ​​the circuit board 10 is clamped between the two clamps 12. Sealing rings 11 are respectively provided between the two clamps 12 and both sides of the circuit board 10. The sealing rings 11 surround the entire nanopore transistor biomolecular sensor. The two clamps 12 are fixed together by bolts on both sides, applying compressive force to the sealing rings 11 and the circuit board 10. The sealing rings 11 are O-rings made of fluororubber.

[0096] Both clamps 12 contain solution pools, with external flow ports and detection flow ports connecting the solution pools. The detection flow port is aligned with the nanopore transistor biomolecular sensor, making the nanopores 9 of the nanopore transistor biomolecular sensor the sole channel connecting the two solution pools.

[0097] In some embodiments, the clamp 12 is made of polytetrafluoroethylene.

[0098] The preparation process of the biomolecule detection device is as follows:

[0099] Step 1: Prepare 10g of uncured PDMS material for later use, and place it in a vacuum environment for 30 minutes to remove air bubbles.

[0100] Step 2: Connect the pads corresponding to the source electrode 4-1 and drain electrode 4-2 to the pads of the interdigitated electrode layer 6 using conductive silver paste and silver wires, respectively, to the corresponding pads on the circuit board 10. Place on a hot plate at 150°C for 10 minutes to solidify the conductive silver paste.

[0101] Step 3: Cover the conductive silver paste portion and the gap between the nanoporous transistor biomolecular sensor and the circuit board 10 with PDMS material. Place it on a hot plate and heat at 100°C for 10 minutes to cure the PDMS, thus isolating the electrode portion from the external environment and sealing the nanoporous transistor biomolecular sensor and the circuit board 10.

[0102] Step 4: Place the circuit board 10 with the sealed nanopore transistor biomolecular sensor on a 60W plasma cleaner for 180 seconds to improve the hydrophilicity of the chip surface.

[0103] Step 5: Place the two sealing rings 11 on the front and back of the circuit board 10, respectively. Embed the circuit board 10 between the two clamps 12 and tighten the clamps 12 with bolts, so that the two clamps 12 clamp the detection area of ​​the circuit board 10. During the clamping process, the two sealing rings 11 are in close contact with the circuit board 10 and the clamps 12, respectively, thus forming a sandwich structure of clamp 12-sealing ring 11-circuit board 10. This achieves front and back wrapping of the circuit board 10, with independent solution pools on both sides of the sensor, and the nanopore 9 becoming the only channel for the solution pools on both sides.

[0104] Example 4

[0105] A biomolecule detection method using the biomolecule detection device provided in Example 3.

[0106] The biomolecule detection method includes the following steps:

[0107] Step 1: Add anhydrous ethanol to the solution pools on both sides of the sensor and heat it on a 50°C hot plate for 10 minutes to allow ethanol molecules to pass through the nanopore 9 and expel air bubbles, thereby improving the wettability of the nanopore 9.

[0108] Step 2: Clean and fill the solution cell with 1M LiCl electrolyte. The LiCl electrolyte is prepared by dissolving LiCl crystals in EDTA (ethylenediaminetetraacetic acid) buffer and keeping the electrolyte at a weakly alkaline pH of 8.

[0109] Step 3: Add two silver-silver chloride reference electrodes to a solution cell filled with LiCl electrolyte and connect them to patch clamps. Determine whether the nanopore 9 is wetted by measuring the current between the two patch clamps. The slope of the nanopore 9 I-V curve is positively correlated with the pore size. If the slope of the recorded IV curve is lower than that of the actual nanopore 9 curve, it indicates poor wettability, and steps 1 and 2 above need to be repeated to improve the hydrophilicity of the nanopore 9. If there is an excessive current, it indicates leakage, and the sealing of the biomolecular detection device needs to be checked and repaired.

[0110] Step 4: Prepare the biomolecule solution. Dilute the biomolecule solution to be tested to 1 nM with LiCl electrolyte and replace the pure LiCl electrolyte in both solution cells.

[0111] Step 5: Apply a constant source-drain voltage (VDS) to the source electrode 4-1 and the drain electrode, and apply a patch-clamp voltage (VG) to adjust the transistor's operating state to the linear region.

[0112] Step 6: Biomolecules (such as DNA) pass through nanopore 9 under the drive of an electric field. The surface charge of the biomolecules disturbs the local electric field of the channel, causing the drain current (ID) to change in real time.

[0113] Step 7: Apply a high-frequency electrical signal to the interdigitated electrode layer 6, and surface acoustic waves are formed due to the piezoelectric effect.

[0114] Step 8: Once the molecular perforation events displayed by the source leakage current reach a relatively stable state, qualitative identification and quantitative detection of biomolecules can be achieved by analyzing the ID fluctuation signal.

[0115] The effect of surface acoustic wave input power on the perforation velocity of biomolecules was tested, and the results are as follows: Figure 7 As shown. The test used a nanopore integrated device with a wavelength of 16μm surface acoustic wave, and the input power was gradually increased from 20dBm to 30dBm.

[0116] Figure 7 Parts a, b, and c respectively show the typical via current signals of the λDNA molecule at input power of 24 dBm, 26 dBm, and 28 dBm. Figure 7 Parts d, e, and f respectively show the through-hole current signals of IgG protein molecules at input power of 24dBm, 26dBm, and 28dBm.

[0117] Test results show that the input power of surface acoustic waves has a significant modulating effect on the molecular through-pore velocity. Specifically:

[0118] Low power range (≤24dBm): The acoustic field effect is weak, and the retardation effect on molecules is not significant. Within this range, the permeation rate of λDNA only shows a slight decrease, such as... Figure 7 As shown in part a, its via time is approximately 0.5 ms.

[0119] In the medium to high power range (≥26dBm and ≤28dBm): the acoustic radiation force and sound field intensity are significantly enhanced, and molecules experience stronger resistance during passage through nanopores, resulting in a significant decrease in the permeation speed. This effect is particularly pronounced for chain-like macromolecules (such as λDNA). At 26dBm power, the permeation time of λDNA is extended to approximately 15ms. Figure 7 (Part b). When the power increases to 28 dBm, the acoustic field hindrance effect reaches its peak, and the perforation time of λDNA is further significantly extended to approximately 45 ms. Figure 7 (Part c). For the relatively compact IgG protein molecule, a similar trend of decreasing velocity was observed, with perforation times of approximately 0.1 ms, 1.5 ms, and 4 ms at 24 dBm, 26 dBm, and 28 dBm powers, respectively.

[0120] Power limit (>28dBm): When the input power exceeds 28dBm, the acoustic device generates significant heat accumulation, resulting in observable evaporation of the electrolyte, and the detection current baseline becomes extremely unstable, making effective measurement impossible.

Claims

1. A nanopore transistor biomolecular sensor based on surface acoustic wave confinement, comprising a substrate (1), a silicon nitride thin film layer (2), a detection layer (4) stacked sequentially, and a nanopore penetrating the sensor; characterized in that: It also includes a first passivation layer (3) covering the detection layer (4), a surface acoustic wave excitation structure disposed on the first passivation layer (3), a second passivation layer (7) covering the surface acoustic wave excitation structure, and a hydrophilic gate medium oxide layer (8) covering the inner peripheral surface of the nanopore. The detection layer (4) includes a source electrode (4-1), a drain electrode (4-2), and a carbon nanotube channel structure (4-3) connecting the source electrode (4-1) and the drain electrode (4-2); the nanopore (9) passes through the carbon nanotube channel structure (4-3); the surface wave excitation structure is used to excite surface acoustic waves focused in the nanopore, including a piezoelectric film (5) and an interdigitated electrode layer (6) disposed on the piezoelectric film (5).

2. The nanoporous transistor biomolecular sensor according to claim 1, characterized in that: The carbon nanotube channel structure (4-3) is made of single-walled semiconductor carbon nanotubes; the side of the carbon nanotube channel structure (4-3) facing away from the silicon nitride thin film layer (2) has an exposure window surrounding the nanopore (9).

3. The nanoporous transistor biomolecular sensor according to claim 1, characterized in that: The interdigitated electrode layer (6) includes two annular interdigitated electrode units symmetrically arranged on both sides of the nanopore (9).

4. The nanoporous transistor biomolecular sensor according to claim 1, characterized in that: The hydrophilic gate medium oxide layer (8) is made of aluminum oxide; the hydrophilic gate medium oxide layer (8) completely covers the cross-section of the first passivation layer (3) and the second passivation layer (7) in the nanopore, and covers the outer surface of the second passivation layer (7).

5. A biomolecule detection method, characterized in that: Using the nanoporous transistor biomolecular sensor according to claim 1; the biomolecular detection method includes: An electrolyte containing the biomolecules to be tested is introduced into a nanopore; a gate voltage is applied to the electrolyte, and a source-drain voltage is applied to the source electrode (4-1) and the drain electrode (4-2); an excitation signal is applied to the interdigitated electrode layer (6), causing the piezoelectric film (5) to generate surface acoustic waves, which produce a blocking effect on the biomolecules passing through the nanopore; the biomolecules to be tested are detected based on the current change signal of the source electrode (4-1) and the drain electrode (4-2).

6. The biomolecule detection method according to claim 5, characterized in that: The input power of the surface acoustic wave is 26dBm to 28dBm.

7. A method for fabricating a sensor, characterized in that: Used to prepare the nanoporous transistor biomolecular sensor as described in claim 1; The sensor manufacturing method includes: Silicon nitride thin films are grown on both sides of a silicon wafer as a substrate (1); Marking points are patterned on the front side of the obtained substrate, and silicon nitride is removed by patterning on the back side to expose the etching window area; The substrate (1) on the back side of the substrate is etched; the etching depth is 80% to 90% of the substrate thickness; The substrate was immersed in a carbon nanotube toluene dispersion, which allowed the carbon nanotubes to be deposited on the substrate surface. After heating and curing the carbon nanotubes, excess carbon nanotubes were etched away to form a carbon nanotube channel structure (4-3). After forming the source electrode (4-1) and drain electrode (4-2) connecting the carbon nanotube channel structure (4-3) on the front side of the substrate, the first passivation layer (3) is covered. Zinc oxide is grown by magnetron sputtering on the first passivation layer (3) to form a piezoelectric thin film (5). After forming an interdigitated electrode layer (6) on the piezoelectric thin film (5), a second passivation layer (7) is covered. Etching removes the first passivation layer (3), the piezoelectric film (5), and the second passivation layer (7) in the area surrounding the nanopore (9); Deep silicon etching was performed on the back side of the substrate (1) to expose a silicon nitride window; After creating the nanopores (9), aluminum oxide is grown on the inner surface of the nanopores (9) as a hydrophilic gate medium oxide layer (8).

8. A biomolecular detection device, characterized in that: The device includes a circuit board (10), a sealing ring (11), and a clamp (12), and also includes a nanoporous transistor biomolecular sensor as described in claim 1; the nanoporous transistor biomolecular sensor is fixed in the detection area of ​​the circuit board (10); the detection area of ​​the circuit board (10) is sandwiched between two clamps (12).

9. A biomolecular detection device according to claim 8, characterized in that: Two clamps (12) are respectively provided with sealing rings (11) between the two sides of the circuit board (10); the sealing rings (11) surround the detection area of ​​the circuit board (10); the two clamps (12) are fixed to each other and pressurize the sealing rings (11).

10. A biomolecular detection device according to claim 8, characterized in that: The circuit board (10) is also provided with an interface area; a detection signal interface (10-1) and an ultrasonic excitation signal interface (10-2) are fixed on the interface area of ​​the circuit board (10); the source electrode (4-1) and drain electrode (4-2) on the nanoporous transistor biomolecular sensor are led out to the detection signal interface (10-1) through the conductive layer on the circuit board (10); the interdigitated electrode layer (6) on the nanoporous transistor biomolecular sensor is led out to the ultrasonic excitation signal interface (10-2) through the conductive layer on the circuit board (10); the conductive layer of the circuit board (10) and the electrical connection structure of the nanoporous transistor biomolecular sensor, as well as the conductive layer within the detection area, are all sealed and hydrophilic.

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