CMOS compatible biosensor based on two-dimensional semiconductor and preparation method and application thereof

Abstract

The invention discloses a CMOS compatible biosensor based on a two-dimensional semiconductor and a preparation method and application of the CMOS compatible biosensor, specifically, an aluminum nitride layer compatible with a CMOS process is deposited on a silicon substrate, a two-dimensional semiconductor layer is formed on the surface of the aluminum nitride layer to serve as a sensing layer, and meanwhile a biological recognition element is fixed. The two-dimensional semiconductor layer is used for being combined with target biomolecules to generate a reaction product with electron donating and receiving capability, so that the reaction product is subjected to efficient charge transfer on the two-dimensional semiconductor layer, and the conductive detection layer can output sensitive and controllable signal change; the band gap characteristic and the high surface area of the two-dimensional semiconductor are utilized, the aluminum nitride and the silicon substrate are combined for use to achieve CMOS compatibility of the biosensing device, and due to the fact that the two-dimensional semiconductor has the excellent optical absorption and emission characteristics, the biosensing device can also be expanded to optical biosensing, optical biosensing and the like. Coupling detection of photoelectric signals and chemical or biological events is achieved, and the application range and multifunctionality are further expanded.

Description

Technical Field

[0001] This invention belongs to the field of micro-nano manufacturing technology, and particularly relates to a CMOS-compatible biosensor based on two-dimensional semiconductors, its preparation method, and its application. Background Technology

[0002] With the development of life sciences and medical diagnostics, highly sensitive, highly selective, and biocompatible sensing technologies are becoming increasingly important in fields such as disease detection, environmental monitoring, and food safety. Biosensors, as key devices combining biometrics and signal conversion, enable rapid and accurate detection of specific biomolecules (such as proteins, nucleic acids, and enzymes).

[0003] At present, the research on biosensors mainly focuses on the following types of channel materials: (1) Traditional semiconductor materials: such as silicon-based FETs, bioMEMS devices, etc. Although their processing technology is mature and compatible with existing CMOS technology, their sensitivity and dynamic range are still limited when detecting low-concentration molecules due to the limited on / off ratio and insufficient surface sensitivity; (2) Carbon-based materials: such as graphene and carbon nanotubes. Graphene has been widely studied due to its high carrier mobility and excellent surface conductivity. However, its zero bandgap characteristic leads to poor field-effect switching performance. The device is difficult to achieve complete turn-off at low carrier concentration, which affects the signal contrast and detection dynamic range of the biosensor. In addition, surface functionalization has a significant impact on device performance and requires complex modification processes; (3) Functionalized nanomaterials: such as quantum dots, nanowires, etc. Although these materials exhibit high sensitivity in specific applications, their large-scale preparation, stability and compatibility with standard semiconductor processes still present challenges, limiting their application in integrated biosensor platforms.

[0004] Overall, while existing biosensor technologies have made significant progress in sensitivity, selectivity, and real-time detection capabilities, they still face challenges such as the trade-off between sensitivity and switching performance, insufficient compatibility with standard semiconductor processes, and limited dynamic response range. Therefore, how to achieve biosensors that are compatible with mature semiconductor processes while maintaining high sensitivity, high selectivity, and wide dynamic range has become a pressing technical problem that needs to be solved in this field. Summary of the Invention

[0005] Based on this, the present invention aims to propose a CMOS-compatible biosensor based on two-dimensional semiconductors, its fabrication method and application. By depositing aluminum nitride on the surface of a silicon substrate as a substrate and using two-dimensional semiconductors as a sensing layer, a wide detection dynamic range and CMOS compatibility can be achieved, thereby realizing wafer-level integrated chip manufacturing.

[0006] In a first aspect, the present invention provides a CMOS-compatible biosensor based on two-dimensional semiconductors, comprising:

[0007] silicon substrate;

[0008] An aluminum nitride layer, as a piezoelectric layer, is deposited on the surface of a silicon substrate in a manner compatible with CMOS processes;

[0009] A conductive detection layer and a two-dimensional semiconductor layer are formed on the surface of the aluminum nitride layer. The two-dimensional semiconductor layer is formed as a sensing layer in a manner compatible with CMOS processes.

[0010] Biometric elements fixed on the surface of a two-dimensional semiconductor layer are used to bind with target biomolecules to generate reaction products with electron donation and acceptance capabilities. The reaction products undergo charge transfer in the two-dimensional semiconductor layer, thereby causing changes in the output signal of the conductive detection layer.

[0011] Furthermore, a buffer layer is formed between the silicon substrate and the aluminum nitride layer.

[0012] Furthermore, the buffer layer comprises silicon dioxide and / or aluminum oxide.

[0013] Furthermore, the aluminum nitride layer is deposited on the silicon substrate surface in a manner compatible with CMOS processes, including:

[0014] The deposition method of aluminum nitride layer formation on silicon substrate surface meets the temperature window of standard CMOS process.

[0015] Furthermore, the aluminum nitride layer has a thickness of 3 μm.

[0016] Furthermore, the aluminum nitride layer is c-axis oriented aluminum nitride.

[0017] Furthermore, the aforementioned biosensor also includes a microfluidic channel for guiding the target biomolecule to bind to the biorecognition element.

[0018] Furthermore, microfluidic channels are etched onto the surface of the aluminum nitride layer.

[0019] Furthermore, a microfluidic channel is disposed on the surface of a microfluidic channel layer, which is disposed above the biometric element and covered by a two-dimensional semiconductor layer.

[0020] Furthermore, the microfluidic channel layer disposed above the biometric element and covering the two-dimensional semiconductor layer includes:

[0021] After being bonded to the aluminum nitride layer, the microfluidic channel layer is located above the biometric element and covers a two-dimensional semiconductor layer.

[0022] Furthermore, the width of the microfluidic channel is 10~100μm and the height is 50~500μm.

[0023] Furthermore, the material of the two-dimensional semiconductor layer includes at least one of transition metal chalcogenides, black phosphorus, chalcogenide-doped composite materials, or a combination of at least one of the above materials.

[0024] Furthermore, the preferred material for the two-dimensional semiconductor layer is molybdenum disulfide.

[0025] Furthermore, the biosensor also includes an acoustic transducer electrode structure for generating surface acoustic waves by voltage triggering, in order to control the flow rate and / or flow direction of target biomolecules in the microfluidic channel.

[0026] Furthermore, the conductive detection layer adopts an interdigitated electrode structure.

[0027] Furthermore, the electrode spacing of the interdigitated electrode structure of the conductive detection layer is no greater than 100 μm, and the effective area is no greater than 1.2 mm × 1.2 mm.

[0028] Furthermore, the acoustic transducer electrode structure includes an interdigitated electrode structure with an electrode spacing of 2~15μm.

[0029] Furthermore, the material of the interdigitated electrode structure includes at least one of gold, silver, copper, aluminum, molybdenum, chromium, nickel, platinum, titanium, and tungsten, or an alloy including at least one of gold, silver, copper, aluminum, molybdenum, chromium, nickel, platinum, titanium, and tungsten.

[0030] In a second aspect, the present invention provides a method for fabricating a CMOS-compatible biosensor based on a two-dimensional semiconductor, applicable to the fabrication of the aforementioned CMOS-compatible biosensor based on a two-dimensional semiconductor, comprising:

[0031] Take a silicon substrate and sputter an aluminum nitride layer on the surface of the silicon substrate using a deposition method compatible with CMOS processes;

[0032] A conductive detection layer and a two-dimensional semiconductor layer are formed on the surface of an aluminum nitride layer, wherein the two-dimensional semiconductor layer is formed as a sensing layer in a manner compatible with CMOS processes.

[0033] Biometric elements are fixed on the surface of a two-dimensional semiconductor layer.

[0034] Furthermore, sputtering an aluminum nitride layer using a deposition method compatible with CMOS processes includes:

[0035] The temperature window of standard CMOS process is met when sputtering and depositing an aluminum nitride layer on a silicon substrate.

[0036] Furthermore, the aluminum nitride layer is deposited by reactive magnetron sputtering deposition, with a deposition temperature not exceeding 400℃ and a post-annealing temperature not exceeding 300℃.

[0037] Furthermore, forming a two-dimensional semiconductor layer on the surface of the aluminum nitride layer includes:

[0038] A two-dimensional semiconductor thin film is prepared on the surface of a growth substrate, and a support layer is coated on the surface of the two-dimensional semiconductor thin film;

[0039] A water stripping method was used to separate a two-dimensional semiconductor thin film coated with a support layer from the growth substrate.

[0040] The two-dimensional semiconductor thin film coated with a support layer, which has been separated from the growth substrate, is transferred to a silicon substrate on which an aluminum nitride layer is formed, so that the two-dimensional semiconductor thin film is formed on the surface of the aluminum nitride layer.

[0041] The two-dimensional semiconductor layer is obtained by removing the support layer on the surface of the two-dimensional semiconductor thin film.

[0042] Furthermore, the above method also includes:

[0043] A buffer layer is deposited between the silicon substrate and the aluminum nitride layer to form a buffer layer.

[0044] Furthermore, depositing a buffer layer between the silicon substrate and the aluminum nitride layer includes:

[0045] A 10 nm thick silicon dioxide layer and / or a 10 nm thick aluminum oxide layer are deposited between a silicon substrate and an aluminum nitride layer.

[0046] Furthermore, the above method also includes:

[0047] Microfluidic channels are set on the aforementioned CMOS-compatible biosensor based on two-dimensional semiconductors to guide the binding of target biomolecules with biorecognition elements.

[0048] Furthermore, setting the microfluidic channel includes:

[0049] Grooves are etched on the surface of the aluminum nitride layer to form microfluidic channels.

[0050] Furthermore, setting the microfluidic channel includes:

[0051] A microfluidic channel layer is prepared, on which a microfluidic channel network is formed;

[0052] A microfluidic channel layer is placed above the biometric element and covered with a two-dimensional semiconductor layer.

[0053] Furthermore, placing the microfluidic channel layer above the biometric element and covering it with a two-dimensional semiconductor layer includes:

[0054] After the microfluidic channel layer is activated, it is bonded to the aluminum nitride layer, so that the microfluidic channel layer is located above the biometric element and covered by a two-dimensional semiconductor layer.

[0055] The present invention also provides an application of the CMOS-compatible biosensor based on two-dimensional semiconductor provided in the first aspect above in the detection of biomolecules based on enzyme reaction products.

[0056] As can be seen from the above technical solution, the present invention has the following beneficial effects:

[0057] This invention provides a CMOS-compatible biosensor based on two-dimensional semiconductors, its fabrication method, and its applications. By depositing an aluminum nitride layer compatible with CMOS processes on a silicon substrate, and forming a two-dimensional semiconductor layer as a sensing layer on its surface, while simultaneously fixing biorecognition elements, efficient charge transfer is achieved between the two-dimensional semiconductor layer and the reaction products of target biomolecules. This enables the conductive detection layer to output sensitive and controllable signal changes. This invention utilizes the bandgap characteristics and high surface area of ​​two-dimensional semiconductors, resulting in a sensor with a high on / off ratio and a wide dynamic response range, improving the detection sensitivity for low concentrations of target biomolecules. Further embodiments employ an interdigitated electrode structure for the conductive detection layer and an acoustic transducer electrode structure, allowing precise control of the flow rate and direction of target molecules in the microfluidic channel, achieving rapid and repeatable biomolecule detection. The design of the microfluidic channel and the two-dimensional semiconductor... The combined coverage of the conductive sensing layer enhances the stability and spatial resolution of the sensing signal. During fabrication, a support layer and water stripping technology are employed to achieve low-temperature transfer of the two-dimensional semiconductor thin film, ensuring the integrity of the two-dimensional semiconductor layer and CMOS compatibility. In a preferred embodiment, the buffer layer effectively reduces stress and interface defects, improving device reliability. The overall solution can be integrated with standard CMOS processes, facilitating large-scale fabrication and industrialization. It is also suitable for highly sensitive detection based on enzyme reaction products or other biomolecules, significantly improving the detection performance, repeatability, and engineering application value of biosensors. Furthermore, due to the excellent optical absorption and emission characteristics of two-dimensional semiconductors, the sensor device provided by this invention can be extended to optical biosensing, enabling the coupled detection of photoelectric signals and chemical or biological events, further expanding its application scope and multifunctionality. Attached Figure Description

[0058] 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 embodiments of the present invention. For those skilled in the art, other drawings can be obtained based on the provided drawings without creative effort.

[0059] Figure 1 This is a schematic diagram of the structure of a CMOS-compatible biosensor based on two-dimensional semiconductors provided in one embodiment of the present invention;

[0060] Figure 2This is a cross-sectional view of a CMOS-compatible biosensor based on two-dimensional semiconductors provided in one embodiment of the present invention;

[0061] Figure 3 yes Figure 2 The schematic diagram shows the functional area of ​​the device, where Figure 3 (a) shows a top-view scanning electron microscope (SEM) image of the sensor device. Figure 3 (b)~3(h) show the EDS energy spectrum diagrams of the material element distribution of the sensor device, which represent the distribution of aluminum (Al), nitrogen (N), oxygen (O), silicon (Si), copper (Cu), molybdenum (Mo), and sulfur (S) in sequence;

[0062] Figure 4 This is a schematic diagram of the Raman peak of a single layer of molybdenum disulfide provided in an embodiment of the present invention;

[0063] Figure 5 This is a graph showing the relationship between the output current response and glucose concentration when an embodiment of the present invention is applied to glucose detection;

[0064] Figure 6 This is a graph showing the relationship between sensitivity and glucose concentration when an embodiment of the present invention is applied to glucose detection;

[0065] Figure 7 This is an illustration of an extended use of a CMOS-compatible biosensor based on two-dimensional semiconductors provided in one embodiment of the present invention. Detailed Implementation

[0066] The technical solutions of the embodiments of the present invention will be clearly and completely described below with reference to the accompanying drawings. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. Based on the embodiments of the present invention, all other embodiments obtained by those skilled in the art without creative effort are within the scope of protection of the present invention.

[0067] The application of two-dimensional materials in the field of biosensors has been extensively studied. Among them, graphene, due to its atomic-level thickness, high carrier mobility, and good surface adsorption capacity, is widely used for the detection of various biomolecules. Graphene can interact with target molecules by immobilizing biorecognition elements, thereby changing their surface charge or carrier concentration and achieving signal output. However, due to graphene's zero-bandgap structure, its field-effect switching performance is poor, and it is prone to response saturation under low concentration or weak signal conditions, limiting its detection capability and signal contrast over a wide dynamic range.

[0068] In contrast, two-dimensional semiconductors possess an atomically layered lattice structure, with the lattice often arranged in a hexagonal or rhombic pattern. Electrons exhibit a two-dimensional quantum confinement effect within the lattice, thus endowing them with a finite band structure and a tunable band gap. This finite band gap allows two-dimensional semiconductors to maintain low leakage current and a high on / off ratio even at low carrier concentrations, significantly improving the sensor's sensitivity to trace biomolecules. Simultaneously, the high specific surface area and surface chemical activity of two-dimensional semiconductors provide abundant fixation sites for biorecognition elements, making charge transfer between biomolecules and the sensing layer more controllable. This broadens the sensor's dynamic response range, achieving highly sensitive and linear detection performance.

[0069] Based on the aforementioned characteristics, two-dimensional semiconductors, as sensing layer materials, offer significant advantages over graphene with its zero bandgap in biosensing applications requiring high on / off ratios and wide dynamic ranges. Therefore, this invention selects two-dimensional semiconductors as the sensing layer material and AlN as the piezoelectric layer material to construct an AlN / Si substrate, and integrates a two-dimensional semiconductor layer on its surface, thereby proposing a CMOS-compatible biosensor based on two-dimensional semiconductors.

[0070] In the embodiments provided below, an AlN / Si substrate is used as the integration basis. Combining the physical and chemical properties of two-dimensional semiconductor materials, since two-dimensional semiconductors can be prepared by processes such as low-temperature deposition, sputtering, or water stripping and transfer, they can be well integrated with aluminum nitride / silicon substrates while maintaining structural integrity and electronic performance. This makes them compatible with standard CMOS processes and suitable as a sensing layer on a conductive detection layer for highly sensitive biomolecule detection.

[0071] See Figure 1 and Figure 2 One embodiment of the present invention provides a CMOS-compatible biosensor based on two-dimensional semiconductors, comprising:

[0072] Silicon substrate 101;

[0073] An aluminum nitride layer 102 is deposited on the surface of a silicon substrate 101 as a piezoelectric layer in a manner compatible with CMOS processes.

[0074] A conductive detection layer and a two-dimensional semiconductor layer 104 are formed on the surface of the aluminum nitride layer 102. The two-dimensional semiconductor layer 104 is formed as a sensing layer in a manner compatible with CMOS process.

[0075] The biometric element 105 fixed on the surface of the two-dimensional semiconductor layer 104 is used to bind with the target biomolecule to generate a reaction product with electron donation and acceptance capabilities. The reaction product undergoes charge transfer in the two-dimensional semiconductor layer 104, thereby causing a change in the output signal of the conductive detection layer.

[0076] In this embodiment, the two-dimensional semiconductor layer is specifically disposed on the surface of the aluminum nitride layer, and may further be disposed on the aluminum nitride layer on which the conductive detection layer has been formed, so that it is stacked with the conductive detection layer, and is used as a sensing layer to sense biomolecular reaction products and cause changes in electrical signals.

[0077] The material for the two-dimensional semiconductor layer can be selected from transition metal chalcogenides (e.g., molybdenum disulfide). ), Tungsten disulfide ( ), molybdenum diselenide ( ), tungsten diselenide ( ), molybdenum ditelluride ( The materials used include single-layer or few-layer materials such as black phosphorus (BP), chalcogenide-doped composites, etc., or any combination of the above materials. These materials can be used alone or in combination to modulate the electronic properties, bandgap characteristics, and interfacial chemical activity of the sensor, meet the detection needs of different biomolecules, and ensure compatibility with CMOS processes.

[0078] In a preferred embodiment, the two-dimensional semiconductor material is molybdenum disulfide (Mo). ). It has a layered lattice structure, with each layer consisting of a molybdenum (Mo) atom layer sandwiched between two sulfur (S) atom layers, forming a sandwich structure. The layers are stacked together by van der Waals forces. This lattice structure endows it with... The limited direct or indirect band gap (approximately 1.2–1.8 eV) in monolayer or few-layer configurations allows for low leakage current and high field-effect on / off ratios at low carrier concentrations, contributing to enhanced sensor sensitivity and linear response range for trace target molecules. Electrons in The layer is constrained by two-dimensional quantum processes, resulting in high mobility. Simultaneously, the film surface possesses abundant active sites, facilitating the immobilization of various biorecognition elements (such as RNA, DNA, and enzymes) and enhancing charge transfer efficiency after binding with target molecules. Its high specific surface area and the chemical activity of the lattice surface not only improve the interaction efficiency between biomolecules and the sensing layer but also significantly broaden the sensor's dynamic response range, enabling the sensor to obtain stable and linear signal output under both low and high concentration conditions.

[0079] also, The light absorption and emission properties of the layer offer potential extensions for optical biosensing, enabling the sensing layer to be used for both electrical detection and multimodal detection in conjunction with optical signals. As a two-dimensional semiconductor layer, this embodiment can achieve high sensitivity, wide dynamic range and functionalizable surface biosensing under the condition of maintaining CMOS compatible process, and provide stable and reliable signal output for target molecule detection.

[0080] In terms of surface coupling performance, two-dimensional semiconductors have ultra-thin atomic layer thicknesses, ranging from a single to several atomic layers, with almost all atoms exposed on the surface. This makes them extremely sensitive to the external environment. Their high specific surface area provides more active sites for the immobilization and binding of biomolecules. Furthermore, the absence of dangling bonds on the surface allows them to interact with other materials or molecules primarily through van der Waals forces, reducing interfacial defects and helping to maintain the controllability and reproducibility of the binding process. When two-dimensional semiconductors are immobilized and coupled with biorecognition elements (such as RNA, DNA, and enzymes), these elements can be strongly and stably immobilized or modified through electrostatic interactions, π–π bond interactions, and edge sites within the two-dimensional semiconductor layer.

[0081] Specifically, in this embodiment of the invention, aluminum nitride (AlN) is selected as the piezoelectric layer. Aluminum nitride (AlN) is a wide bandgap semiconductor with a hexagonal wurtzite structure. Its c-axis direction has excellent piezoelectric coupling coefficient and high sound velocity, which is suitable for the generation and control of surface acoustic waves or leakage surface acoustic waves. Therefore, AlN thin films can be directly deposited on silicon substrate 101 by reactive magnetron sputtering to further form a highly c-axis oriented (002) structure, so that the piezoelectric performance can be fully utilized, and at the same time, it is compatible with the substrate system commonly used in silicon-based integrated circuits.

[0082] Traditional piezoelectric materials such as lithium niobate, lead zirconate titanate (PZT), and zinc oxide (ZnO) often require process temperatures exceeding 500°C during deposition or annealing, or contain elements like lead that are unsuitable for CMOS process lines. In contrast, AlN thin films can be deposited at temperatures below 400°C, with post-annealing temperatures not exceeding 300°C, fully complying with the temperature limits of standard CMOS back-end metal interconnects (BEOL). This avoids damage to fragile layers such as copper interconnects and low-k dielectrics, enabling the biosensor to be directly embedded into silicon-based CMOS processes, making mass production feasible.

[0083] Furthermore, AlN possesses chemical inertness and excellent surface stability, making it resistant to corrosion or degradation in complex biofluid environments. It also contains no toxic elements such as lead or cadmium, meeting the safety requirements for biomedical applications. Therefore, using aluminum nitride as the piezoelectric layer in this invention simultaneously ensures device stability and biofriendliness.

[0084] Through the structural design of this invention, the aluminum nitride layer 102 not only provides the piezoelectric function basis for the sensor, but also makes the entire structure compatible with silicon-based CMOS processes, ensuring that the subsequent construction of conductive detection layer 103, two-dimensional semiconductor layer 104, etc. can be realized under standard process conditions. As a result, the biosensor proposed in this invention can be manufactured with low cost, high yield and high integration using silicon process platform, while maintaining excellent sensing sensitivity and stability.

[0085] In a further embodiment, c-axis oriented (002) aluminum nitride is selected for the aluminum nitride layer. This is because AlN has a hexagonal wurtzite structure, and its piezoelectric properties are closely related to its crystal orientation. Specifically, c-axis (002) oriented AlN films have the highest longitudinal piezoelectric coefficient and strong coupled piezoelectric response. Their orientation is characterized by X-ray diffraction (XRD), which shows a significant (002) diffraction peak at 2θ≈36.0°, with the (002) peak intensity accounting for ≥80% and a full width at half maximum (FWHM) ≤2°. These parameters indicate that the film has good crystal quality and is highly c-axis oriented, making it a preferred orientation for realizing surface acoustic wave (SAW) and bulk acoustic wave (BAW) devices. Compared with non-oriented or randomly oriented AlN films, c-axis oriented films have better lattice order and electrical homogeneity, which is beneficial for reducing noise and improving signal output stability. In biosensors, c-axis oriented AlN layers can generate a stronger surface electric field modulation effect, thereby improving the charge control sensitivity of the two-dimensional semiconductor layer.

[0086] In another embodiment, the aluminum nitride layer thickness is selected to be approximately 3 μm. This thickness ensures sufficient piezoelectric coupling strength, while process optimization avoids the generation of stress cracks, allowing the structure to still maintain CMOS process compatibility.

[0087] Specifically, depositing aluminum nitride on a silicon substrate in a manner compatible with CMOS processes allows the deposition method to meet the temperature window of standard CMOS processes. In advanced CMOS integrated circuit manufacturing, back-end interconnect (BEOL) processes typically require deposition and post-processing temperatures not exceeding 400°C, while post-annealing or surface repair processes are usually limited to below 300°C. This temperature window limitation is primarily to avoid adverse effects on the formed metal interconnects, low-dielectric-constant insulating layers, and doped profiles. Therefore, by utilizing the temperature window to limit the deposition process of the aluminum nitride layer, the resulting aluminum nitride layer 102 not only possesses excellent piezoelectric properties and low-stress characteristics but also does not compromise the material integrity and electrical properties required by silicon-based CMOS processes.

[0088] In a preferred embodiment, the deposition temperature of the aluminum nitride layer is ≤400℃, and the post-annealing temperature is ≤300℃; reactive magnetron sputtering is employed. A more preferred embodiment can further control parameters such as the nitrogen partial pressure ratio range, sputtering RF power, deposition rate, and dielectric constant during aluminum nitride layer deposition, enabling the prepared aluminum nitride layer to exist stably and function within a standard CMOS process system, thereby achieving a truly CMOS-compatible biosensor.

[0089] In a further embodiment, due to the significant differences in lattice constants and coefficients of thermal expansion between silicon and aluminum nitride, directly depositing an aluminum nitride layer on the silicon substrate surface can easily lead to stress concentration during deposition or subsequent heat treatment, causing the aluminum nitride layer to crack or peel off, thereby affecting device performance and reliability. By introducing a buffer layer between silicon and aluminum nitride, lattice mismatch and thermal expansion mismatch can be effectively alleviated, improving the crystal quality and c-axis orientation of the aluminum nitride film, enhancing interfacial adhesion, reducing internal stress in the film, and providing a smoother surface for subsequent deposition of two-dimensional semiconductor layers and electrodes.

[0090] Specifically, Figure 1 A buffer layer 106 is schematically disposed between a silicon substrate 101 and an aluminum nitride layer 102. The buffer layer may be made of silicon dioxide. ) layer, aluminum oxide ( ) layer, silicon nitride ( (Layer and other materials or combinations thereof)

[0091] In another embodiment, the buffer layer is composed of and Layered construction, for example, first forming a 10nm thick layer on a silicon substrate. Then deposit a 10nm thick layer This balances interfacial adhesion and film stability.

[0092] In the biometric identification section, the biometric element used in this embodiment needs to undergo a specific reaction after binding with the target biomolecule to produce reaction products with electron donation and acceptance capabilities, such as hydrogen peroxide (H2O2), metal ions, charged metabolites, nitrogen oxides, etc. Since the two-dimensional semiconductor layer is highly sensitive to electron injection and depletion, these interfacial process changes are converted into measurable conductivity changes under a fixed bias voltage, thereby reflecting the target biomolecule and its concentration changes through the electrical signal output by the sensor.

[0093] In a further embodiment, the conductive detection layer 103 can adopt an interdigitated electrode structure (IDE). Due to the interlocking and nesting of the electrode interdigitates, a large-area lateral electric field distribution is formed, which can create multiple current paths. Utilizing this structural feature, the effective interface area between the two-dimensional semiconductor layer or the fluid to be measured can be increased, thereby improving the conductivity response sensitivity and comprehensively acquiring changes in the conductivity of the two-dimensional semiconductor layer. The electric field lines of the IDE are mainly distributed horizontally between the electrode interdigitates, making it more suitable for the detection of two-dimensional materials or thin films. Furthermore, compared to the two-end electrode structure, the current channel density per unit area of ​​the IDE is higher, making it particularly suitable for low-concentration detection or small-signal acquisition scenarios, resulting in a more sensitive sensor response in low-concentration or low-dose detection tasks. The interdigitation gap can be defined as the deposition area of ​​the functional film layer. For example, when preparing a two-dimensional semiconductor material film layer on it, the effective area of ​​the two-dimensional semiconductor material can be precisely controlled, improving selectivity or stability. In addition, the multi-finger structure of the IDE can disperse local noise interference, making it easier to average the micro-area response, especially when biomolecules are unevenly distributed or react unevenly, making it more stable.

[0094] Furthermore, as a compact device design, the interdigitated electrode structure of the conductive detection layer has an electrode spacing of no more than 100 μm, an interdigitated electrode length of no less than 500 μm, and an effective area of ​​no more than 1.2 mm × 1.2 mm. It does not require a large area of ​​material, and the sensor device has a compact structure. A two-dimensional semiconductor layer can be formed by transferring a two-dimensional semiconductor thin film onto the interdigitated electrode structure.

[0095] For example, Figure 1 The illustration shows a representative design where the interdigitated electrode structure of the conductive detection layer has an electrode length of 800 μm and the conductive detection area is only 1000 μm * 1000 μm in size, which can minimize the amount of material used while maintaining stable conductivity.

[0096] like Figure 2 As shown, it illustrates a cross-sectional view of a biosensor with a biometric element fixed thereon. From bottom to top, the elements are a silicon substrate 201, a buffer layer 202, an aluminum nitride layer 203, a conductive detection layer 204, a molybdenum disulfide layer 205, and a biometric element 206.

[0097] Specifically, Figure 2 The conductive detection layer 204 includes two electrodes that are separated from each other, serving as source 2041 and drain 2042 respectively. A molybdenum disulfide layer 205 covers the conductive detection layer 204 and forms a conductive channel between the source and drain.

[0098] During operation, the source 2041 serves as the current injection terminal, and the drain 2042 serves as the current output terminal. The current between them is conducted through the two-dimensional semiconductor layer 205. When the biorecognition element 206 fixed on the surface of the molybdenum disulfide layer 205 specifically binds to the target biomolecule, a charge transfer effect or local potential change occurs on the surface of the molybdenum disulfide layer, thereby modulating the carrier concentration in the molybdenum disulfide channel and causing a change in the conductivity between the source and drain. By measuring the current change between the source and drain, highly sensitive detection of the target biomolecule can be achieved.

[0099] Unlike conventional methods that only use electrodes as impedance testing terminals, this embodiment introduces a source / drain structure in the conductive detection layer 204, enabling the sensor of this invention to operate similarly to a two-dimensional semiconductor field-effect transistor, thus demonstrating its semiconductor device characteristics. This design not only improves signal sensitivity and stability but also makes the sensor easier to integrate with CMOS circuit structures, thereby facilitating mass production, low-cost fabrication, and on-chip system applications.

[0100] like Figure 3 As shown in (a)~3(h), Figure 3 (a) shows a top-view scanning electron microscope (SEM) image of the sensor device. Figure 3 (b)~3(h) show the EDS energy spectrum diagrams of the material element distribution of the sensor device, which sequentially represent the distribution of aluminum (Al), nitrogen (N), oxygen (O), silicon (Si), copper (Cu), molybdenum (Mo), and sulfur (S), confirming the stacked state of each layer before the biometric element is fixed.

[0101] For example, Figure 4 It also indicated a single layer A type of Raman peak, in which E 2g The peak is 378cm -1 Location, A 1g The peak is at 404cm -1 Place.

[0102] Based on the foregoing embodiments, the working process of the CMOS-compatible biosensor based on two-dimensional semiconductors provided in this invention mainly includes the following:

[0103] The target biomolecule is dropped onto the biorecognition element of the biosensor. The biorecognition element reacts with the target biomolecule to produce reaction products with electron donation and acceptance capabilities. The reaction products inject or deplete electrons into the two-dimensional semiconductor layer located below the biorecognition element. That is, the Fermi level of the two-dimensional semiconductor is changed through charge transfer or redox events, which changes the conductivity of the two-dimensional semiconductor and thus changes the output current of the biosensor.

[0104] For example, taking glucose oxidase as a biorecognition element, the application process includes:

[0105] A glucose sample solution is dropped onto the biosensor's biorecognition element, where it binds to glucose oxidase. Simultaneously, a bias voltage (e.g., 3.0V) is applied to the interdigitated electrodes to detect the electrical signal. Glucose oxidase catalyzes glucose, converting it into gluconic acid, accompanied by the generation of hydrogen peroxide (H2O2). Each H2O2 molecule injects two electrons into the two-dimensional semiconductor layer, thereby increasing the Fermi level of the two-dimensional semiconductor. The increased conductivity of the two-dimensional semiconductor leads to an increase in the output current. Therefore, the output current changes before and after the glucose is dropped.

[0106] In a further embodiment, the biometric element fixed on the biosensor provided by the present invention can be selected from, for example, oxidases, oxorubicins, nitric oxide-related enzymes, polyphenolases, antibodies, aptamers, molecularly imprinted polymers, etc.

[0107] The working mechanism of the biosensor provided by the present invention will be further illustrated by a specific example below.

[0108] Specifically, the sensor uses a high-resistivity silicon substrate, and a 10nm thick layer is deposited on the surface of a 300μm silicon substrate. and 10nm thick As a buffer layer, a 3 μm thick aluminum nitride layer was deposited on the surface of the buffer layer by reactive magnetron sputtering at a deposition temperature not exceeding 400 °C and a post-annealing temperature not exceeding 300 °C to improve crystallinity and piezoelectric response. Afterwards, it was washed with deionized water and dehydrated at 120 °C for 10 minutes. The interdigitated electrode structure pattern of the conductive detection layer was fabricated using optical lithography, with a distance of 100 μm between the interdigitated electrodes. Sputtering deposition was used, with Cr (10 nm) as the adhesive layer and Au (50 nm) as the contact layer. The layer was stripped with acetone, rinsed with IPA (isopropanol), and dried in nitrogen (N2) to obtain the patterned conductive detection layer. The layer was then deposited on a growth substrate (e.g., Monolayer on / Si) The substrate is coated with PMMA (polymethyl methacrylate) as a support layer and separated using a water stripping method. and growth substrate, to obtain Thin film, will The thin film is transferred to the device region (e.g., AlN / Si), then the PMMA layer is removed using acetone and IPA (isopropanol), and dried with nitrogen (N2) to form a film. Layer; glucose oxidase as a biorecognition element, glucose oxidase solution prepared in buffer solution (>pH5), spin-coated onto... The surface of the layer is then freeze-dried at -40°C.

[0109] A small droplet containing the target analyte is introduced into the sensing area. A 3.0V bias voltage is applied across the conductive detection layer, and the converted signal is monitored as the source-drain current. Taking the detection of glucose using glucose oxidase as an example, glucose is oxidized to gluconic acid, and simultaneously... ;Reaction products mediate charge direction Layer transfer, thereby altering the Fermi level and increasing Electrical conductivity. Therefore, with increasing glucose concentration (e.g., 5, 10, and...),... The Fermi level will shift, which corresponds to a gradual increase in the output current of the sensor. Therefore, the concentration of the target analyte can be quantitatively analyzed by the change in the output current of the device before and after its introduction.

[0110] The output current of the device is detected, and as the enzymatic reaction proceeds, the reaction products move towards... Electrons are injected into the layer to increase its Fermi level and increase the carrier density, thereby increasing the output current of the device under a fixed bias voltage. The current change is proportional to the concentration of the target biomolecule. Table 1 below further illustrates the detection results of the output current of the biosensor provided in this embodiment, which shows the output current change corresponding to different glucose concentrations (0~30mmol / L).

[0111] Table 1

[0112]

[0113] The data in Table 1 show that the device output current increases steadily with glucose concentration. When the concentration changes from 0 to... At the same time, the measured current also increases proportionally; with appropriate calibration at a bias voltage of 3.0V, the average sensitivity is approximately . Figure 5 and Figure 6 The current-glucose concentration characteristic curve of the output current signal and the sensitivity change curve based on the data in Table 1 were plotted.

[0114] In a further embodiment, the biosensor also includes a microfluidic channel for guiding the target biomolecule to bind to the biorecognition element.

[0115] Microfluidic channels can guide target biomolecules to flow along a predetermined path, enabling efficient sample transport and reaction. In conjunction with aluminum nitride piezoelectric layers and acoustic transducer electrode structures, they can achieve localized pumping, mixing, and flow direction control of fluids (e.g., using surface acoustic waves to drive liquids). At the same time, they provide protection for biometric elements on the surface of two-dimensional semiconductor layers and prevent liquid leakage and contamination of other functional areas of the device.

[0116] Specifically, in conjunction with the use of microfluidic channels, acoustic transducer electrode structures can also be integrated into the biosensors provided in the aforementioned embodiments. Specifically, it can be an IDE. Utilizing the triggering principle of surface acoustic waves, as long as the interdigitated electrodes are deposited on the surface of the piezoelectric material film, an electrical signal applied to the IDE will excite mechanical deformation and generate surface acoustic waves, which will propagate along the surface of the piezoelectric material film.

[0117] More specifically, it utilizes the leakage surface acoustic wave effect to generate surface acoustic waves through an acoustic transducer electrode structure. Once a liquid (water, buffer solution, cell suspension, etc.) is added to the microfluidic channel, the acoustic impedance matching and critical angle propagation conditions are met at the solid-liquid interface, and part of the energy of the surface acoustic wave will "leak" into the liquid, forming a leakage surface acoustic wave. This enables the device to perform fluid control such as molecular pumping more precisely.

[0118] In one embodiment, the acoustic transducer electrode structure (used to generate surface acoustic waves and control fluid movement within the microfluidic channel) is electrically completely separated from the conductive detection layer.

[0119] In some embodiments, trenches can be directly etched into the aluminum nitride layer to serve as microfluidic channels.

[0120] Specifically, a trench structure with a predetermined width and depth is etched on the surface of an aluminum nitride layer using photolithography and dry etching (e.g., reactive ion etching, RIE) processes. This trench can be closed in subsequent processes by covering it with a transparent encapsulation layer (e.g., PDMS or a glass sheet), thereby forming a closed microfluidic channel.

[0121] The trenches are directly etched onto the surface of the aluminum nitride layer, reducing the need for additional steps involving a separate microfluidic substrate and bonding processes, resulting in a more compact overall device structure. As a piezoelectric material, aluminum nitride, combined with the trenches, can form a surface acoustic wave propagation path near the channel, thereby enabling pumping and stirring of the fluid within the channel when triggered by the acoustic transducer electrode, further improving the binding efficiency of biomolecules and detection sensitivity. The method of etching directly onto the piezoelectric layer still maintains compatibility with CMOS processes, as the trench etching process can be integrated with standard semiconductor micromachining procedures, making it suitable for mass production on silicon substrates.

[0122] In different embodiments, the width of the groove can be designed to be about 10 to 100 μm and the height to be about 50 to 500 μm to meet the channel requirements of different types of biological fluid samples (such as serum, saliva or buffer).

[0123] In a further embodiment, microfluidics can also be achieved by separately fabricated microfluidic channel layers and enclosed devices, with the microfluidic channel layers disposed above the biometric element and covered by a two-dimensional semiconductor layer.

[0124] Specifically, the microfluidic channel layer can be fabricated using PDMS soft lithography or other polymer microfabrication methods to form a network of microfluidic channels of predetermined shape and size. The microfluidic channel layer is then placed on top of the biometric element and covered with a two-dimensional semiconductor layer, thereby forming a closed microfluidic cavity above the conductive detection layer.

[0125] In different embodiments, the microfluidic channel layer and the aluminum nitride layer can be stably encapsulated through physical pressing or chemical bonding (such as bonding after oxygen plasma surface activation), thereby forming a reliable and reusable microfluidic structure. The channel size can be designed according to specific application requirements, for example, a width of about 10~100μm and a height of about 50~500μm, to balance fluid resistance and reaction efficiency.

[0126] Specifically, the microfluidic channel layer is typically prepared using PDMS or other polymeric materials. Its surface has low surface energy, making it prone to debonding or leakage when in direct contact with the aluminum nitride layer. Surface activation treatments, such as oxygen plasma treatment or ultraviolet ozone treatment, can generate hydroxyl (–OH) or oxide groups on the PDMS surface, making it hydrophilic and increasing its chemical reactivity with the aluminum nitride surface. In a more preferred embodiment, oxygen plasma activation treatment can also slightly etch the PDMS surface, creating a nanoscale rough structure, further increasing the interfacial contact area and thus improving physical adhesion.

[0127] By setting up a microfluidic channel layer, this invention can realize on-chip microfluidic functions while maintaining CMOS compatibility, thereby improving the sensitivity, repeatability and ease of operation of biomolecule detection.

[0128] In a further embodiment, the CMOS-compatible biosensor provided by the present invention can achieve multiplexing, that is, multiple independent sensor units are integrated simultaneously on the same silicon substrate. Each unit includes an aluminum nitride layer, a two-dimensional semiconductor layer, and surface-functionalized biometric elements. Each sensor unit can independently measure the concentration of a specific biomolecule, and multiple targets can be detected simultaneously on the same chip, realizing parallel biological analysis. Monolithic integration helps to reduce the overall device size, improve sample utilization, and reduce operational and wiring complexity.

[0129] To achieve monolithic CMOS control, a through-hole electrode can be set in the aluminum nitride layer of each sensor unit. The through-hole directly penetrates the aluminum nitride layer, interconnecting the conductive detection layer or acoustic transducer electrode structure with the CMOS drive / readout circuit on the silicon substrate. Through the through-hole electrode, the CMOS chip can independently perform bias control and current readout for each sensor unit, realizing time-division or parallel acquisition of electronic signals. Thus, this invention realizes a limited sensing to a monolithically integrated sensor chip platform.

[0130] By integrating multiple sensor units on a silicon substrate and directly interconnecting them with CMOS circuits using aluminum nitride through-hole electrodes, a monolithic, multiplexed, high-throughput biomolecule detection scheme can be realized, while maintaining CMOS compatibility and highly integrated system-on-chip functionality.

[0131] Figure 7 The illustration depicts a device expansion using microfluidic channels. Specifically, it illustrates a detection device structure integrating five biosensors. The sensor devices and acoustic transducer electrode structures share the same Al / Si layer as a common substrate. The common substrate includes a stacked silicon substrate, a buffer layer, and an aluminum nitride layer. The microfluidic channels are etched onto a PDMS layer, which is disposed above the sensor devices and covers the two-dimensional semiconductor layers of each device.

[0132] Specifically, five sensors simultaneously acquire data from the same sample to check the accuracy of results in actual detection applications. Specifically, a sample droplet of the target biomolecule can be introduced at 701, moving to the microfluidic channel via capillary effect. The acoustic transducer electrode structure 702 generates surface acoustic waves under the triggering of an external voltage source 703, pumping the sample to five sensing regions S1~S5 respectively. The change in conductivity of the two-dimensional semiconductor layer of each sensor is converted into an electrical signal and sent to electrode 704 (…). Figure 7 (Taking the electrode of one of its sensor devices as an example), this signal current can be further used to analyze the concentration of target biomolecules. Still using glucose detection as an example, Table 2 below illustrates the use of... Figure 7 The schematic diagram shows the results of concentration detection using an extended sensor structure.

[0133] Table 2

[0134]

[0135] By designing the sensor to support on-chip redundancy, self-verification, and simultaneous multi-analyte expansion (e.g., assigning different biometric elements to different channels or sensor locations), while utilizing the same AlN / Si microfluidic sensing platform to achieve a compact, low-power operating system.

[0136] In a further embodiment, a frequency-dependent surface acoustic wave contrast mechanism can be used to guide specific biomolecules to a designated microfluidic channel, thereby enabling the simultaneous detection of different target biomolecules.

[0137] Furthermore, the material of the interdigitated electrode includes at least one of gold, silver, copper, aluminum, molybdenum, chromium, nickel, platinum, titanium, and tungsten, or an alloy including at least one of gold, silver, copper, aluminum, molybdenum, chromium, nickel, platinum, titanium, and tungsten.

[0138] The following embodiments of the present invention also provide a method for fabricating a CMOS-compatible biosensor based on two-dimensional semiconductors.

[0139] Specifically, one embodiment provides a method for fabricating a CMOS-compatible biosensor based on a two-dimensional semiconductor, comprising the following steps:

[0140] Step S101. Take a silicon substrate and sputter an aluminum nitride layer on the surface of the silicon substrate using a deposition method compatible with CMOS process.

[0141] This step involves directly depositing an aluminum nitride layer on the surface of a silicon substrate. By controlling the substrate surface state, buffer layer properties, and deposition parameters (temperature, atmosphere, power, bias voltage, etc.), optimal orientation, low internal stress, and high uniformity of AlN are achieved. As a piezoelectric layer, the deposition process of the aluminum nitride layer must meet the compatibility requirements of CMOS processes, thereby ensuring that the sensor can be manufactured monolithically with silicon-based integrated circuits.

[0142] Silicon, as the standard substrate for CMOS processes, has mature wafer-level fabrication processes and interconnect systems. After aluminum nitride layers are directly sputtered and deposited on the surface of silicon substrates, their strong piezoelectricity, low dielectric loss, and preferred orientation crystal structure can provide piezoelectric support for subsequent electroacoustic coupling and two-dimensional semiconductor electrical detection.

[0143] In some embodiments, the silicon substrate can be a single-crystal silicon wafer, preferably a (100) crystal orientation silicon wafer, to facilitate subsequent micro-nano fabrication processes. The diameter of the silicon wafer can be selected from 200 to 300 mm to meet the compatibility requirements with mainstream CMOS production line processes.

[0144] In some embodiments, a buffer layer may be pre-grown on the surface of the silicon substrate. This buffer layer alters the chemical and atomic arrangement of the interface, providing a better nucleation platform, reducing mismatch dislocation formation, and mitigating residual stress caused by thermal expansion differences. For example, an insulating buffer layer ( , It can also act as a diffusion barrier to protect the underlying metal / device structure.

[0145] In some embodiments, a 10 nm thick layer is deposited between the silicon substrate and the aluminum nitride layer. Layers and / or 10nm thickness layer.

[0146] To ensure compatibility between sensor devices and CMOS processes, the aluminum nitride layer needs to be deposited using a CMOS-compatible deposition method. Specifically, the deposition process of the aluminum nitride thin film on the silicon substrate must meet the overall compatibility requirements of CMOS manufacturing processes to avoid damage or contamination to front-end devices and back-end interconnect structures. For example, the deposition temperature window, material selection, atmosphere conditions, process parameters, and film formation characteristics must all meet the requirements of standard CMOS manufacturing processes.

[0147] Specifically, an aluminum nitride layer is deposited on the silicon substrate using a method compatible with CMOS processes, such as reactive magnetron sputtering, RF magnetron sputtering, metal-organic chemical vapor deposition, or atomic layer deposition. During deposition, the temperature is controlled within a CMOS-acceptable window (e.g., ≤400°C), while the atmosphere (Ar / ) is adjusted. The flow rate ratio, chamber pressure, and sputtering power are adjusted to obtain a high-quality aluminum nitride layer preferentially grown on the c-axis oriented (002) crystal plane; low-temperature annealing can be performed after deposition to reduce stress and improve crystal quality, but not exceeding the thermal budget allowed by the CMOS process.

[0148] In some embodiments, the aluminum nitride layer is deposited by reactive magnetron sputtering at a deposition temperature not exceeding 400°C, preferably 200–350°C, to avoid damaging the formed CMOS circuit, and the post-annealing temperature not exceeding 300°C to release stress without affecting the interconnect metal layer.

[0149] In a preferred embodiment, the aluminum nitride layer is c-axis oriented (002) aluminum nitride, which means that the intensity of the (002) peak in its X-ray diffraction spectrum accounts for more than 90% of the main peak and the half-peak width is less than 2°, so as to ensure the high quality factor of the surface acoustic wave device.

[0150] In some embodiments, the aluminum nitride layer thickness is selected to be about 3 μm to achieve a trade-off between piezoelectric effect and mechanical stability, while reducing the risk of stress cracking.

[0151] After obtaining the silicon substrate (AlN / Si) with an aluminum nitride layer deposited on the silicon substrate surface, further processing is required to improve the cleanliness of the substrate surface, improve the adhesion of subsequent processes, and ensure device consistency, including surface cleaning, drying and dehydration.

[0152] In some embodiments, to improve the adhesion of subsequent films or materials (such as electrodes, two-dimensional semiconductor films, PDMS layers), the surface of the aluminum nitride layer can be activated using oxygen plasma or argon plasma.

[0153] Post-processing can effectively remove residual particles, impurities, and moisture from the sputtering deposition process, improving the cleanliness and flatness of the AlN layer surface. It avoids problems such as poor photoresist adhesion, two-dimensional semiconductor transfer failure, or weak microfluidic layer bonding caused by contamination or water film. Furthermore, this step ensures the CMOS process compatibility and mass production consistency of the device.

[0154] For example, the AlN / Si substrate is cleaned and rinsed with deionized water, and then dehydrated at 120°C for 10 minutes to obtain the piezoelectric layer.

[0155] Step S102. A conductive detection layer and a two-dimensional semiconductor layer are formed on the surface of the aluminum nitride layer, wherein the two-dimensional semiconductor layer is formed as a sensing layer in a manner compatible with CMOS processes.

[0156] In this step, a two-dimensional semiconductor layer can be formed on the surface of the aluminum nitride layer using various processes. Considering CMOS process compatibility, the fabrication of the two-dimensional semiconductor layer should meet the following requirements: firstly, the fabrication temperature should not exceed the temperature window of the CMOS process (generally <400℃) to avoid damaging the underlying metal interconnects and dielectric structures; secondly, the continuity and uniformity of the two-dimensional semiconductor layer should be ensured, and it should be controllable and patterned using standard microfabrication processes.

[0157] Two-dimensional semiconductor layers can be fabricated using chemical vapor deposition (CVD), physical vapor deposition (PVD), stripping transfer, and molecular beam epitaxy. Considering CMOS process compatibility requirements, PVD is often used in conjunction with a transfer process.

[0158] In a preferred embodiment, the transfer of the two-dimensional semiconductor thin film is achieved using a water stripping method. The formation of the two-dimensional semiconductor layer on the surface of the aluminum nitride layer involves the following process:

[0159] A two-dimensional semiconductor thin film is prepared on the surface of a growth substrate, and a support layer is coated on the surface of the two-dimensional semiconductor thin film. The two-dimensional semiconductor thin film coated with the support layer is separated from the growth substrate by a water stripping method. The two-dimensional semiconductor thin film coated with the support layer, which has been separated from the growth substrate, is transferred to a silicon substrate with an aluminum nitride layer, so that the two-dimensional semiconductor thin film is formed on the surface of the aluminum nitride layer. The support layer on the surface of the two-dimensional semiconductor thin film is removed to obtain the two-dimensional semiconductor layer.

[0160] In this embodiment, a water stripping method is preferably used to transfer the two-dimensional semiconductor film on the growth substrate to a silicon substrate with an aluminum nitride layer already formed. Specifically, a two-dimensional semiconductor film is first prepared on the growth substrate, and a polymethyl methacrylate (PMMA) support layer is spin-coated onto the film surface (e.g., 4000 rpm, 60 s, baking at 100°C for 1 min, PMMA thickness approximately 100–300 nm). Then, the PMMA-coated sample is slowly immersed in deionized water, with water penetrating along the substrate / film interface, gradually reducing the interfacial adhesion, causing the PMMA / two-dimensional semiconductor film to separate from the growth substrate and float on the water surface. Subsequently, the target substrate is slowly lifted from below the water to contact and adhere to the floating PMMA / film. After adhesion, the PMMA is dissolved with acetone (e.g., soaked at room temperature for 10–30 min) and rinsed with IPA. Finally, it is annealed at ≤300°C under vacuum or a gas-forming atmosphere to remove organic residues and improve interfacial contact. Water stripping is a gentle process that avoids corrosion of the aluminum nitride layer and metal interconnects by strong acids and alkalis, making it easy to integrate with standard CMOS processes. At the same time, the integrity and electrical / chemical properties of the transferred two-dimensional film can be optimized by controlling the PMMA thickness, stripping speed, bonding strategy and annealing conditions.

[0161] In a preferred embodiment, the growth substrate for the two-dimensional semiconductor can be... , / Si、 Quartz glass and metal foils (such as Cu, Ni, and Mo foils). A more preferred embodiment may use a composite substrate, such as... / Si、 / Si, taking advantage of the weak bonding between the composite substrate and the two-dimensional semiconductor thin film, is more suitable for low-destructive transfer using water stripping method. It is also highly compatible with the temperature window and cleanliness requirements of CMOS process, and can realize the growth of large-area, highly uniform two-dimensional semiconductor thin films, meeting the requirements of biosensor devices for material quality and process compatibility.

[0162] In a preferred embodiment, the conductive detection layer formed on the surface of the aluminum nitride layer is an electrode structure used to extract the electrical signals generated by the two-dimensional semiconductor layer and convert them into detectable voltage or current signals.

[0163] In some embodiments, the conductive detection layer may be in the form of a source and drain pair, and more preferably, the electrode action area may be patterned into an interdigitated electrode structure with a spacing between 1 μm and 100 μm to improve electric field uniformity and detection sensitivity.

[0164] Furthermore, the process of forming interdigitated electrodes commonly uses metallic materials including at least one of gold, silver, copper, aluminum, molybdenum, chromium, nickel, platinum, titanium, and tungsten, or alloys including at least one of gold, silver, copper, aluminum, molybdenum, chromium, nickel, platinum, titanium, and tungsten.

[0165] In some embodiments of the present invention, when the conductive detection layer adopts a source and drain structure, the two-dimensional semiconductor layer is connected between the source and drain, or covers the interdigitated electrode gap, so that the electrode and the two-dimensional semiconductor form an ohmic contact.

[0166] Step S103. Fix biometric elements on the surface of the two-dimensional semiconductor layer.

[0167] The biometric element fixed in this step refers to a functionalized molecule or complex that can specifically bind to a specific target molecule (such as DNA, RNA, protein, antibody, antigen, extracellular vesicle, viral particle, etc.). Specifically, the biometric element can be fixed on a two-dimensional semiconductor layer by drop casting or spin coating and incubated at -40°C for freeze drying.

[0168] In one embodiment, the biometric element covers the entire two-dimensional semiconductor layer to achieve full-surface detection; or, only the effective detection area between the electrodes is locally modified to reduce non-specific adsorption.

[0169] In some embodiments, prior to fixing the biometric element, the non-sensor regions of the two-dimensional semiconductor layer and the conductive detection layer are passivated to reduce non-specific adsorption, lower background signals, and improve detection selectivity. Here, "non-sensor regions" refer to areas in the two-dimensional semiconductor layer or conductive detection layer that do not directly participate in the recognition and signal transduction of target biomolecules. Passivation can be achieved through covering, coating, or chemical modification, causing these regions to lose or reduce their non-specific binding ability to biomolecules, thereby ensuring that the signal source is primarily concentrated in the functionalized sensing active region.

[0170] In a further embodiment, the fabrication process also includes forming microfluidic channels on the device, specifically by etching directly onto the surface of the aluminum nitride layer, or by providing a microfluidic channel layer to define the flow path of the target biomolecules and guide the target biomolecules to bind to the biorecognition element.

[0171] In one embodiment, the microfluidic channel can be formed by selectively etching an aluminum nitride piezoelectric layer deposited on a silicon substrate. The aluminum nitride layer, as a hard material, can be etched into trench structures with regular geometric shapes using reactive ion etching (RIE), inductively coupled plasma etching (ICP), or wet etching methods. Specifically, photoresist can be coated and patterned on the surface of the aluminum nitride layer, using… Plasma gas is used to etch the exposed area, and the etching depth can be controlled in the range of hundreds of nanometers to tens of micrometers.

[0172] In another embodiment, the microfluidic channel is formed by an independent channel layer, preferably polydimethylsiloxane (PDMS) material. The PDMS is pre-formed into the desired channel structure on the SU-8 master mold by soft photolithography, and then bonded to the aluminum nitride layer after activation treatment.

[0173] Specifically, an SU-8 master mold can be prepared, with microfluidic channels of approximately 10-100 μm in width and 50-500 μm in height. PDMS is then poured into the mold and cured to obtain a microfluidic channel layer with grooves. Further oxygen plasma treatment is used to increase the hydroxyl density of PDMS and the AlN surface, thereby achieving bonding. The PDMS layer is then applied over a two-dimensional semiconductor layer on top of the biometric element, defining the liquid flow path.

[0174] Furthermore, in conjunction with the use of microfluidic channels, acoustic transducer electrode structures can be integrated on the surface of the aluminum nitride layer, specifically an IDE structure.

[0175] In some embodiments, the acoustic transducer electrode structure can be formed on the surface of an aluminum nitride layer etched outside a microfluidic channel trench to excite surface acoustic waves within the trench. Since the conductive detection layer is simultaneously located on the two-dimensional semiconductor surface at the bottom or sidewall of the trench, to ensure electrical separation, the acoustic transducer electrode and the conductive detection layer electrode need to be staggered in planar arrangement, or separated by an insulating layer (such as...). Electrical isolation is performed.

[0176] In another embodiment, the acoustic transducer electrode is typically disposed on the surface of the aluminum nitride layer, below or adjacent to the PDMS channel. Since PDMS is an electrically insulating material, the transducer electrode can be located directly below the PDMS channel without contacting the fluid. In this case, the electrode structure of the conductive detection layer and the transducer electrode are electrically separated through planar partitioning.

[0177] For example, the following describes the fabrication process of a CMOS-compatible biosensor based on two-dimensional semiconductors, using glucose oxidase as a biorecognition element:

[0178] Using a high-resistivity silicon substrate, a 10nm thick layer was deposited on the surface of a 300μm thick silicon substrate. and 10nm thick As a buffer layer, a 3 μm thick aluminum nitride layer was deposited on the surface of the buffer layer by reactive magnetron sputtering at a deposition temperature not exceeding 400℃ and a post-annealing temperature not exceeding 300℃ to improve crystallinity and piezoelectric response. Afterwards, it was washed with deionized water and dehydrated at 120℃ for 10 minutes. The interdigitated electrode structures of the conductive detection layer and the acoustic transducer electrode structure were fabricated using optical lithography. The distance between the interdigitated electrodes of the conductive detection layer was 100 μm, and the spacing between the acoustic transducer electrodes was 2–15 μm. The electrode structures of both parts were deposited by sputtering, using Cr (10 nm) as the adhesive layer and Au (50 nm) as the contact layer. The layers were stripped with acetone, rinsed with IPA (isopropanol), and dried in nitrogen (N2). The deposited layers were then deposited on the growth substrate. Monolayer on Si The substrate is coated with PMMA (polymethyl methacrylate) as a support layer and separated using a water stripping method. and growth substrate, to obtain Thin film, will The thin film was transferred to an AlN / Si substrate, then the PMMA layer was removed using acetone and IPA (isopropanol), and dried with nitrogen (N2) to form the final product. Layer; glucose oxidase as a biorecognition element, glucose oxidase solution prepared in buffer solution (>pH5), spin-coated onto... The surface of the layer is then freeze-dried at -40°C; a PDMS (polydimethylsiloxane) film is prepared on the AlN / Si substrate, and a microfluidic channel network is formed on the PDMS layer. The microfluidic channel network is located on the side where the PDMS layer contacts the aluminum nitride layer, and the contact surface between the PDMS layer and the aluminum nitride layer is activated by oxygen plasma, so that the PDMS and the surface of the AlN layer are bonded.

[0179] All embodiments provided in this invention can be fabricated based on standard microelectromechanical systems (MEMS) and CMOS processes. The selected device structure materials, such as piezoelectric materials (e.g., aluminum nitride), conductive electrode materials (e.g., gold, platinum, chromium, ITO), and functional film materials (e.g., PDMS), are all commonly used material systems supported by current MEMS manufacturing platforms, exhibiting good process compatibility and structural repeatability. The structure of each device layer can be achieved through standard MEMS / CMOS processes such as thin film deposition, photolithography, etching, spin coating, surface modification, and bonding packaging, without deviating from the temperature and material compatibility requirements of existing integrated circuit manufacturing processes.

[0180] The sensor devices and microfluidic structures provided by this invention are directly compatible with silicon-based CMOS integrated circuits, enabling monolithic integration, mass production, and high-throughput fabrication. This ensures that the devices achieve industrial-grade performance, reliability, and process consistency. Furthermore, the material selection, microstructure dimensions, channel arrangement, and electrode configuration involved in the various embodiments of this invention can be flexibly adjusted according to specific detection requirements without affecting their manufacturability based on standard MEMS / CMOS processes.

[0181] Therefore, the sensor device described in this invention can be automated and mass-produced using existing MEMS / CMOS processing platforms, possessing excellent process stability and yield control capabilities. It is easy to integrate into existing health management infrastructures such as medical testing or wearable devices. At the same time, this manufacturing path has good adaptability to the existing industrial chain, which helps to achieve low-cost, large-scale mass production of the device in medical, diagnostic, and biological monitoring scenarios.

[0182] It should be understood that the biosensor provided in the embodiments of the present invention is not limited to using two-dimensional semiconductor materials as the electrical sensing layer to achieve detection functions. Since two-dimensional semiconductor materials typically possess tunable band gaps, strong light absorption, and significant photoluminescence characteristics, within the framework of the present invention, the biosensor can also be extended to optical biosensing applications. By combining the optical response characteristics of two-dimensional semiconductors, detection of biological substances based on light signals can be achieved, thereby further improving the applicability and detection accuracy of the sensor of the present invention. Those skilled in the art will understand that the above-described extension to optical biosensing does not deviate from the core concept of the present invention.

[0183] The invention has been described in particular detail above with respect to possible scenarios, and those skilled in the art will recognize that the invention can be practiced through other embodiments. Specific naming of components, capitalization of terms, attributes, data structures, or any other programming or structural aspects are not mandatory or important, and the mechanisms or features of implementing the invention may have different names, forms, or procedures. The system can be implemented through a combination of hardware and software (as described), entirely through hardware elements, or entirely through software elements. The specific division of functions among the various system components described herein is merely exemplary and not mandatory; rather, the functions performed by a single system component can be performed by multiple components, or the functions performed by multiple components can be performed by a single component.

[0184] Those skilled in the art should understand that the various steps of the disclosed methods can be implemented using general-purpose computing devices. They can be centralized on a single computing device or distributed across a network of multiple computing devices. Optionally, they can be implemented using device-executable program code, which can then be stored in a storage device for execution by the computing device. Alternatively, they can be fabricated as separate integrated circuit modules, or multiple modules or steps can be fabricated as a single integrated circuit module. Therefore, the embodiments disclosed in this invention are not limited to any specific hardware and software combination.

[0185] Those skilled in the art will understand that the structures shown in the figures are merely block diagrams of some structures related to the present application and do not constitute a limitation on the terminal device to which the present application is applied. Specific terminal devices may include more or fewer components than those shown in the figures, or combine certain components, or have different component arrangements.

[0186] In the description of this specification, the use of terms such as "one embodiment," "some embodiments," "example," "specific example," or "possible design," etc., refers to a specific feature, structure, material, or characteristic described in connection with that embodiment or example, which is included in at least one embodiment or example of this application. In this specification, the illustrative expressions of the above terms do not necessarily refer to the same embodiment or example. Furthermore, the specific features, structures, materials, or characteristics described may be combined in a suitable manner in any one or more embodiments or examples. Moreover, without contradiction, those skilled in the art can combine and integrate the different embodiments or examples described in this specification, as well as the features of different embodiments or examples.

[0187] When using expressions such as "at least one of A, B, and C", they should generally be interpreted in accordance with the meaning that is commonly understood by a person skilled in the art (e.g., "a system having at least one of A, B, and C" should include, but is not limited to, a system having A alone, a system having B alone, a system having C alone, a system having A and B, a system having A and C, a system having B and C, and / or a system having A, B, and C, etc.).

[0188] The above 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.

Claims

1. A CMOS-compatible biosensor based on two-dimensional semiconductors, characterized in that, include: silicon substrate; An aluminum nitride layer, as a piezoelectric layer, is deposited on the surface of the silicon substrate in a manner compatible with CMOS processes; A conductive detection layer and a two-dimensional semiconductor layer are formed on the surface of the aluminum nitride layer, wherein the two-dimensional semiconductor layer is formed as a sensing layer in a manner compatible with CMOS processes. A biometric element fixed to the surface of the two-dimensional semiconductor layer is used to bind with a target biomolecule to generate a reaction product with electron donation and acceptance capabilities. The reaction product undergoes charge transfer in the two-dimensional semiconductor layer, thereby causing a change in the output signal of the conductive detection layer.

2. The CMOS-compatible biosensor based on two-dimensional semiconductors according to claim 1, characterized in that, The aluminum nitride layer is deposited on the silicon substrate surface in a manner compatible with CMOS processes, including: The aluminum nitride layer is deposited on the surface of the silicon substrate in a manner that meets the temperature window of standard CMOS processes.

3. The CMOS-compatible biosensor based on two-dimensional semiconductors according to claim 1, characterized in that, The sensor also includes a microfluidic channel for guiding the target biomolecule to bind to the biorecognition element.

4. The CMOS-compatible biosensor based on two-dimensional semiconductors according to claim 3, characterized in that, The microfluidic channel is disposed on a microfluidic channel layer, which is disposed above the biometric element and covers the two-dimensional semiconductor layer.

5. The CMOS-compatible biosensor based on two-dimensional semiconductors according to claim 1, characterized in that, The material of the two-dimensional semiconductor layer includes at least one of transition metal chalcogenides, black phosphorus, and chalcogenide-doped composite materials, or a combination of at least one of the materials.

6. The CMOS-compatible biosensor based on two-dimensional semiconductors according to claim 1 or 5, characterized in that, The material of the two-dimensional semiconductor layer includes molybdenum disulfide.

7. A method for fabricating a CMOS-compatible biosensor based on a two-dimensional semiconductor, applied to the fabrication of a CMOS-compatible biosensor based on a two-dimensional semiconductor as described in any one of claims 1 to 6, comprising: Take a silicon substrate and sputter an aluminum nitride layer on the surface of the silicon substrate using a deposition method compatible with CMOS processes; A conductive detection layer and a two-dimensional semiconductor layer are formed on the surface of the aluminum nitride layer, wherein the two-dimensional semiconductor layer is formed as a sensing layer in a manner compatible with CMOS processes. Biometric elements are fixed on the surface of the two-dimensional semiconductor layer.

8. The method according to claim 7, characterized in that, Forming the two-dimensional semiconductor layer on the surface of the aluminum nitride layer includes: A two-dimensional semiconductor thin film is prepared on the surface of a growth substrate, and a support layer is coated on the surface of the two-dimensional semiconductor thin film. The two-dimensional semiconductor thin film coated with the support layer is separated from the growth substrate using a water stripping method. The two-dimensional semiconductor film coated with a support layer, which has been separated from the growth substrate, is transferred to the silicon substrate on which an aluminum nitride layer is formed, so that the two-dimensional semiconductor film is formed on the surface of the aluminum nitride layer. The two-dimensional semiconductor layer is obtained by removing the support layer from the surface of the two-dimensional semiconductor thin film.

9. The method according to claim 7, characterized in that, The method further includes: A microfluidic channel layer is prepared, on which a microfluidic channel network is formed; After the microfluidic channel layer is activated, it is bonded to the aluminum nitride layer, so that the microfluidic channel layer is located above the biometric element and covers the two-dimensional semiconductor layer.

10. The application of a CMOS-compatible biosensor based on two-dimensional semiconductors as described in any one of claims 1 to 6 in the detection of biomolecules based on enzyme reaction products.

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