Microfluidic chip, biomaterial-plastic composite material device using microfluidic chip, and non-contact interconnection mode of microfluidic chip and connector

Abstract

A microfluidic chip, a biomaterial-plastic composite material device using the microfluidic chip, and a non-contact interconnection mode of the microfluidic chip and a connector. The microfluidic chip not only enhances the mechanical strength of biomaterial chips, but also enjoys the advantage of a cell-on-top configuration, thereby ensuring the authenticity of test results. Even in the presence of gaps, a pinning effect can still be used to assist in the connection between the microfluidic chip and the connector, achieving a non-contact interconnection strategy.

Description

Microfluidic chip, biomaterial-plastic composite device using the same, and contactless interconnection method between microfluidic chip and connector Technical Field The present application belongs to the field of microfluidics and biomaterials, relates to a microfluidic chip, a biomaterial-plastic composite material device using the same, and a non-contact interconnection method between the microfluidic chip and a connector. Background Art Microfluidic chips are a type of technology platform that can integrate or substantially integrate basic operating units such as sample preparation, reaction, separation, detection, cell culture, sorting and lysis involved in fields such as chemistry and biology onto a substrate of several square centimeters or even smaller. A network of microchannels and reaction chambers is formed with controllable fluids running through the entire system, replacing various functions of conventional chemical or biological laboratories. It has the advantages of flexible combination, large-scale integration, miniaturization, automation, integration and high throughput, and is an important means of automation, miniaturization and portability of biological detection equipment. Biomaterials such as hydrogels are emerging as important materials for constructing microfluidic chips. They possess high permeability to small molecules, excellent breathability, and outstanding biocompatibility and biodegradability. Hydrogels are extremely hydrophilic polymer networks with a high water content (20%-90%) and a three-dimensional (3D) porous structure. Due to their extremely high water content, hydrogels can closely mimic the properties of the extracellular matrix and simulate the in vivo microenvironment. By combining hydrogels with microfluidics, hydrogel-based microfluidic chips can simultaneously possess the perfusion capabilities of microfluidics and the free diffusion capacity of hydrogels. Therefore, these chips allow for the spatiotemporal control of multiple signals. Due to the irreplaceable properties of biomaterial-based microfluidic chips, they are now widely used in large-scale in vitro tissue or organ models, as well as in vivo simulations of physiology and pathology, and for drug screening. Summary of the Invention However, the brittleness and softness of biomaterials present obstacles to the assembly and manipulation of connectors. Unlike traditional microfluidic chips, hydrogels cannot be connected using self-elastic forces, connectors, or adhesives. These issues not only create incompatibility between hydrogel-based microfluidic chips and traditional connection processes, but also limit their flexibility in design and functionality. Currently, common connection methods for biomaterial-based microfluidic chips can be categorized by their operation mode, such as insertion, crimping, and integration. While insertion and crimping methods are simple, they can easily damage the chip or cause channel collapse during the connection process, especially for microfluidic channels with high aspect ratios. Therefore, these methods require highly trained personnel and can be very time-consuming if the chip design is complex. Furthermore, crimping methods require custom fixtures, which increases the cost of changing chip designs. In contrast, integration methods eliminate the need for additional fixtures for insertion or crimping and can utilize conventional chips as housings for automated connection to biomaterial chips. Integrated designs typically embed biomaterials into conventional chips through filling or 3D bioprinting. However, when filling specific chambers with hydrogels, more complex chips (such as those with micropillars or encapsulated porous membranes) must be used to block the liquid hydrogel to prevent leakage before solidification. Using 3D printing technology to embed hydrogels requires a cumbersome manufacturing process. First, the sacrificial microfluidic channel must be printed into a Luer adapter for connection. Two hydrogel building blocks are required to support the sacrificial layer, and after cross-linking, the sacrificial microfluidic channel must be cleaned to construct the microfluidic channel. Therefore, the integration method is more expensive and complex than the insertion and press-fit methods. Furthermore, because the press-fit and integration methods use a fixture or traditional chip to clamp or embed the biomaterial, experimental reactions cannot be performed on the surface of the biomaterial. This deprives biomaterial-based microfluidic chips of their advantage in specifically simulating in vivo conditions, such as the inability to simulate biofilm growth, and experimental results may be biased by the influence of fluid shear forces. Therefore, current connection methods still cannot provide a satisfactory connection solution for biomaterial-based microfluidic chips. Means for solving problems This application is completed in view of the above-mentioned actual situation and provides a microfluidic chip, a biomaterial-plastic composite material device using the same, and a non-contact interconnection method between the microfluidic chip and a connector. The purposes are as follows. (1) Improve the mechanical strength of biomaterial chips to achieve automation and large-scale production. (2) Reduce the difficulty of connecting the microfluidic chip and the connector. (3) No additional connection components are required. (4) Provides guidance for inlet and outlet design and allows for gaps to exist and have a wider acceptable range. This application is made through the following methods [1]~

[0011] And completed. Method [1] A microfluidic chip, which is a composite material chip obtained by arranging a biomaterial on a plastic substrate, wherein: A microfluidic channel is constructed on the plastic substrate. Method [2] A microfluidic chip as described in method [1], wherein the biomaterial is polyacrylamide hydrogel, or a combination of polyacrylamide hydrogel and natural materials or peptides. Method [3] The microfluidic chip as described in method [2], wherein the biomaterial is polyacrylamide hydrogel. Method [4] A microfluidic chip as described in method [2] or [3], wherein the natural material is agar or agarose or alginate or collagen. Method [5] The microfluidic chip as described in method [1], wherein the plastic substrate is formed of polymethyl methacrylate, polystyrene, polypropylene, or polyethylene. Method [6] The microfluidic chip as described in method [5], wherein the plastic substrate is formed of polymethyl methacrylate. Mode [7] A biomaterial-plastic composite material device, which is composed of the microfluidic chip and connector according to any one of modes [1] to [6], wherein: The microfluidic channel on the plastic substrate has ports serving as an inlet and an outlet, respectively. The channel on the connector has ports serving as an inlet and an outlet, The fluid flows in the biomaterial-plastic composite material device in the order of the outlet of the connector, the inlet of the microfluidic channel, the outlet of the microfluidic channel, and the inlet of the connector. The surface of the connector is formed with a Teflon coating, There is a gap between the microfluidic chip and the connector, that is, the connection between the two is non-contact. Mode [8] The biomaterial-plastic composite material device according to Mode [7], wherein: The distance of the gap is 60 to 335 μm. Mode [9] The biomaterial-plastic composite material device according to Mode [8], wherein: The distance of the gap is 60 μm. Mode

[0010] The biomaterial-plastic composite material device according to Mode [7], wherein: The radius of the outlet of the connector is smaller than the radius of the inlet of the microfluidic channel, The radius of the inlet of the connector is larger than the radius of the outlet of the microfluidic channel. Method

[0011] A non-contact interconnection method between a microfluidic chip and a connector, wherein: The microfluidic chip is a microfluidic chip according to any one of the methods [1] to [6], The microfluidic channel on the plastic substrate of the microfluidic chip has ports serving as an inlet and an outlet, respectively. The channel on the connector has ports serving as an inlet and an outlet, The fluid flows in the biomaterial-plastic composite material device in the order of the outlet of the connector, the inlet of the microfluidic channel, the outlet of the microfluidic channel, and the inlet of the connector. The surface of the connector is formed with a Teflon coating, There is a gap between the microfluidic chip and the connector. Effects of the Invention The microfluidic chip of the present application protects the biomaterial through a plastic substrate and reduces the difficulty of connection, so as to prevent damage to the biomaterial during operation and realize automation. In addition, the microfluidic chip is used and interconnected with a connector having a Teflon coating in a contact manner. This new interconnection strategy provides another way to connect microfluidic chips without the need for additional connection components or integrated chip design. Therefore, users can make full use of the functions of the microfluidic chip to conduct tests to obtain real results. Compared with other non-contact connection strategies used in traditional microfluidic chips, the interconnection strategy of the present application also provides irreplaceable value for the biomaterial chip connection theory. BRIEF DESCRIPTION OF THE DRAWINGS Figure 1a shows a schematic diagram of a hydrogel-plastic composite device. Figure 1b shows the hydrogel-plastic composite chip being pushed directly into the Teflon-coated (black) connector for connection. Figure 1c illustrates the bonding procedure for the hydrogel-plastic composite chip. Specifically, an open microstructure is constructed on PMMA by CNC, which is then immersed in a 10% benzophenone solution for coating. A physically cross-linked hydrogel is placed on top of the treated PMMA, forming a closed microstructure. After 5 minutes of UV irradiation, the hydrogel undergoes chemical cross-linking, forming a PAAm hydrogel that adheres to the PMMA. The composite chip is then immersed in water to remove unreacted reactants and finally immersed in the target medium. Figure 1d shows a photograph of a hydrogel-plastic composite chip. Figure 1e shows a hydrogel-plastic composite device perfused with red dye. Figure 2ai shows the fluid flow direction of the biomaterial-plastic composite device. Specifically, a peristaltic pump first perfuses the fluid into the chip connector. After passing through the gap, the fluid flows into the composite chip and follows the geometry of the microfluidic channels constructed on the PMMA. Figure 2aii shows a microfluidic channel covered with biomaterial. Small molecules in the fluid diffuse into the biomaterial simultaneously. Thus, the chip can generate a concentration gradient across the biomaterial surface through diffusion. Finally, the fluid flows out of the chip and returns to the chip connector as waste. Figure 2aiii shows the roughness of the Teflon coating. Figure 2bi shows that the fluid is fixed in the gap, forming a liquid bridge, rather than leaking, due to the surface properties of the Teflon coating. A 2D image of the gap during the primary perfusion process is captured using confocal microscopy using fluorescent dyes. FIG2bii shows a 3D image of the interstitial space during primary perfusion, taken by confocal microscopy using fluorescent dyes. Figure 2ci shows a schematic diagram of the leak prevention principle for a contactless connection at the inlet, where the upstream port radius is significantly larger than the downstream port radius. r represents the inlet and outlet radius, α is the PMMA port opening angle, θ is the contact angle, Z is the gap distance, ΔP is the Young-Laplace pressure, subscript P represents PMMA, subscript T represents Teflon, and subscript PT represents the contact interface between PMMA and Teflon. FIG2cii shows a schematic diagram of the anti-leakage principle in which the radius of the upstream port is significantly smaller than that of the downstream port in the case of contactless connection at the inlet. FIG2di shows a schematic diagram of the anti-leakage principle when the outlet is in a contactless connection state and the radius of the upstream port is significantly larger than that of the downstream port. FIG2dii shows a schematic diagram of the anti-leakage principle when the outlet is in a contactless connection condition and the radius of the upstream port is significantly smaller than that of the downstream port. Figure 3 shows that the principle of suspended microfluidics enables stacking and connecting layers without bonding and without leakage outside the lateral channels. Figure 4a illustrates the leak-proof principle using parallel superhydrophobic coatings, showing liquid flowing from one microfluidic chip to another; the capillary force caused by the superhydrophobic coating prevents the lateral expansion of the liquid bridge. Figure 4b shows the leakage prevention using a parallel superhydrophobic coating. Figure 4c shows the leakage situation without the parallel superhydrophobic coating. Figure 5a shows the static pressure diagram of the measured water flow inflow (from Teflon to PMMA) at different gap distances (60-600 μm). Figure 5b shows the static pressure diagram of the measured water outflow (from PMMA to Teflon) at different gap distances (60-600 μm). Figure 6a shows the 2D AST results of a drug combination study on a composite chip using E. coli expressing green fluorescent protein as the sample. The ampicillin gradient was run from right to left, while the gentamicin gradient was generated from top to bottom on the device. FIG6 b shows the growth morphology of E. coli at a high gentamicin concentration and a low ampicillin concentration. FIG6 c shows the growth morphology of E. coli at a medium gentamicin concentration and a low ampicillin concentration. FIG6 d shows the growth morphology of E. coli at a medium gentamicin concentration and a high ampicillin concentration. FIG. 6e shows the growth morphology of E. coli at low concentrations of gentamicin and ampicillin. FIG6 f shows the growth morphology of E. coli at a gentamicin concentration slightly lower than the MIC value and a low ampicillin concentration. FIG6g shows the growth morphology of E. coli at a medium gentamicin concentration and a medium ampicillin concentration. FIG6h shows the growth morphology of E. coli at low gentamicin concentration and high ampicillin concentration. Explanation of symbols 1…Drug storage area, connected to the peristaltic pump 2…Temperature controller 3…Cultivation chamber, where the mixing connector is placed and used to add bacterial samples 4…Central control panel for controlling drug flow rate and incubation time (heating stops and cooling begins after the timer expires) 5…Hydrogel 6…Composite material chips 7…Teflon coating 8…Aluminum 9…PMMA 10…Superhydrophobic coating DETAILED DESCRIPTION The following describes the embodiments of the present disclosure with reference to the accompanying drawings, etc. However, the present disclosure can be implemented in a variety of different ways and is not to be construed as being limited to the description of the embodiments illustrated below. In addition, in order to explain more clearly, the drawings sometimes schematically represent the width, thickness, shape, etc. of each part compared to the actual method, but this is always an example and does not limit the interpretation of the present disclosure. In addition, in this specification and the drawings, the same symbols are given to the elements that are the same as the above-mentioned elements in the drawings that have already appeared, and the detailed description is sometimes appropriately omitted. It should be noted that in this specification, "microfluidic chip" is sometimes specifically referred to as "hydrogel-plastic composite chip," which has the same meaning. Furthermore, PMMA stands for polymethyl methacrylate, PAAm stands for polyacrylamide, and Teflon stands for Teflon. <Microfluidic Chip> The microfluidic chip of the present application is a composite material chip obtained by arranging biomaterials on a plastic substrate, and a microfluidic channel is constructed on the plastic substrate. As a column, a microfluidic chip obtained by the following manufacturing method can be cited. First, an open microfluidic channel is constructed on PMMA by methods such as numerical control, hot stamping or laser cutting, and then the surface is treated with a benzophenone solution. The solidified hydrogel is then directly covered on the microfluidic channel to form a closed channel. After irradiation with ultraviolet light for 5 minutes, the hydrogel undergoes chemical cross-linking to form polyacrylamide (PAAm), and at the same time chemically bonds to the PMMA surface to adhere to the PMMA. Finally, all unreacted reactants in the hydrogel are removed by soaking to ensure the quality of the microfluidic chip. The bonding strategy proposed in this application is shown in Figure 1b. The biomaterial is a polyacrylamide hydrogel, or a combination of a polyacrylamide hydrogel and a natural material or peptide, preferably a polyacrylamide hydrogel. The natural material is agar, agarose, alginate, or collagen. The plastic substrate is formed of polymethyl methacrylate, polystyrene, polypropylene, or polyethylene, preferably polymethyl methacrylate. <Biomaterial-Plastic Composite Device> The biomaterial-plastic composite device of the present application is composed of the microfluidic chip and connector of the present application, wherein the microfluidic channel on the plastic substrate of the microfluidic chip has ports serving as an inlet and an outlet, respectively, and the channel on the connector has ports serving as an inlet and an outlet, respectively. The fluid flows in the biomaterial-plastic composite device in the order of the outlet of the connector, the inlet of the microfluidic channel, the outlet of the microfluidic channel, and the inlet of the connector. A Teflon coating is formed on the surface of the connector. A gap is present between the microfluidic chip and the connector, that is, the two are connected in a non-contact manner. The biomaterial-plastic composite device connects the composite chip based on the interfacial surface tension difference between the plastic substrate (eg, made of polymethyl methacrylate) of the microfluidic chip and the Teflon layer of the connector. The connector is processed from aluminum (Al) or the like, and a Teflon coating is formed thereon by spraying and sintering, thereby increasing the hydrophobicity of the contact surface between the connector and the chip. As described above, the microfluidic channel on the plastic substrate of the microfluidic chip has ports serving as an inlet and an outlet, respectively, and the channel on the connector has ports serving as an inlet and an outlet, respectively. Since the plastic substrate of the microfluidic chip and the ports of the connector have a special geometric design, the chip (Fig. 1c) will automatically align with and connect to the chip port after the connector is pushed in (Fig. 1d). Further, preferably, the radius of the outlet of the connector is smaller than the radius of the inlet of the microfluidic channel, and the radius of the inlet of the connector is larger than the radius of the outlet of the microfluidic channel. The reason is that this design not only makes the connection stronger, but also allows a certain degree of misalignment. This is because the upstream channel is always covered by the downstream channel. Therefore, this design not only reduces the manufacturing difficulty of maintaining the gap distance, but also reduces the complexity of precise microscale alignment. Thus, the biomaterial-plastic composite material device of the present application does not need to be connected to the chip by traditional connection methods. The device is ready for use and can be operated automatically without the need for trained personnel. Specifically, in terms of connection, the connector first applies pressure to the plastic substrate (i.e. PMMA) with a spring, bringing the microfluidic chip as close to the connector as possible. During the liquid flow process, the hydrophobic Teflon coating repels the liquid passing through the connection and forces it into the The hydrophilic PMMA channel is used to complete the connection (Figure 2ai). It should be noted that if there is no Teflon coating, since both Al and PMMA are hydrophilic, the liquid will infiltrate the connection and eventually cause leakage. At the same time, since the Teflon coating has roughness (Figure 2aiii), there will be a certain degree of gap between the chip and the connector, so the two are not in direct contact (Figure 2bi, Figure 2bii). It can be seen that the connection between the microfluidic chip and the connector in this application is a non-contact interconnection strategy. On the other hand, since the liquid is not injected into the hydrogel, the concentration gradient of the liquid is only generated by diffusion in the hydrogel. This situation can make the test unaffected by shear force to ensure the authenticity of the results obtained. <Contactless connection between microfluidic chip and connector> The contactless interconnection mode of the microfluidic chip of the present application and the connector is characterized in that the microfluidic chip and the connector of the present application are used. In this way, the microfluidic chip can still be connected to the connector using the pinning effect assistance even when there is a gap, thus realizing a contactless interconnection strategy. The contactless interconnection mode of the microfluidic chip of the present application and the connector is simple, does not require accurate manufacturing process or alignment operation (allowing connection with gaps), and can be mass-produced. Therefore, the bottleneck of hydrogel chip connection is overcome, automation and large-scale production are realized, and the microfluidic chip is more suitable for clinical use. In order to more clearly understand the technical significance of the contactless interconnection method of the present invention, a conventional connection method is given below for illustration. Conventional connection strategies use either a unique port design (hereinafter referred to as the first conventional method) or a superhydrophobic interface (hereinafter referred to as the second conventional method) to connect conventional chips. In the first traditional method, a double inverted eaves design (T-shape) needs to be manufactured at the microscale at the port to maintain the fluid through the surface tension between the gaps of two vertically suspended microfluidic channels. Due to the Laplace pressure, the fluid in the gap is not easy to leak and will form a curved surface, and then flow into the channel. This T-shaped design is a key element of the leakage prevention strategy. Without a T-shaped design to pin the fluid, the gap used to connect the two closely contacting devices fails to seal, forming a new suspended microfluidic channel (a in Figure 3). Eventually, the connection will become a new lateral channel, and leakage will occur in the gap. In the second traditional method, parallel super-hydrophobic surfaces are used to form an energy barrier to achieve connection. A super-hydrophobic coating (for example, silica particles mixed with epoxy resin) must be applied on both sides of the interconnected surfaces, which are called parallel super-hydrophobic surfaces. When the liquid passes through, the liquid can form a liquid bridge at the connection without leaking. An axisymmetric liquid bridge is formed at the connection and then fixed to the edge of the through-hole by the super-hydrophobic surface. The researchers showed that the use of parallel super-hydrophobic surfaces is crucial to preventing leakage, otherwise leakage will always occur (Figure 4a, Figure 4b, Figure 4c). In addition, the parallel super-hydrophobic surfaces need to be smooth and have a constant contact angle, and the port through-holes need to be the same. This is because a smooth connection is one of the necessary conditions for the stability of the liquid bridge. However, experimental results by the applicants of this application have shown that leak prevention requires neither a T-shaped port design nor parallel, smooth, superhydrophobic surfaces with identical through-holes. The applicants have demonstrated that ports with flat surfaces can still be connected, and that Laplace pressure can be generated by surface tension differences rather than by a double-inverted eaves structure. Furthermore, parallel superhydrophobic surfaces are not essential for connection. Due to a certain degree of surface roughness and surface tension differences, a rough hydrophobic and hydrophilic interface can also secure liquids at the connection. Unlike previous conventional methods, since different surfaces have different contact angles, an asymmetric liquid bridge will be formed instead of an axisymmetric one. In addition, since the liquid is pinned by the rough hydrophobic surface, its stability does not rely on the superhydrophobicity provided by the smooth surface. Therefore, the hydrophobic surface does not need to be smooth. Therefore, due to the roughness provided by the hydrophobic surface, the liquid is allowed to be pinned at the edge of the port, and then the liquid will further pin on the downstream channel surface (rather than just at the edge). Therefore, ports with different radii are acceptable. <Pinning Effect in the Contactless Interconnection Method of the Present Application> According to Laplace's law, fluids tend to shrink spherically to minimize potential energy. From this perspective, if the fluid can maintain a curved liquid surface in the gap (cohesive force dominates), it means that it is in a stable state and the chip can be successfully connected to the connector. By comparing the Young-Laplace pressure (curvature pressure) between the port and the gap, the applicant can estimate whether the fluid is connected or leaking at the connection. The applicant assumes that the radius of the inlet and outlet ports is larger or smaller than the other (there are no identical holes); the inner walls of the PMMA substrate, Teflon coating and aluminum port are rough (the port is drilled); the Teflon coating is rough and has a fixed thickness; the gravity effect can be ignored; and the fluid is water. In general, according to the Young-Laplace equation, the capillary pressure difference (ΔP) is proportional to the contact angle, ΔP = -2γcosθ / r, where γ is the interfacial tension, θ is the contact angle, and r is the port radius. However, at the connection point, the surfaces of two different materials will cause the fluid to form an asymmetric curved shape. Therefore, it is necessary to take into account the surface tension of the two materials, ΔP_PT = -[4γ(cosθ_P+cosθ_T)] / Z, where Z is the gap distance, subscript P is PMMA, subscript T is Teflon, and subscript PT is the connection interface between PMMA and Teflon. If ΔP_PT is greater than ΔP of the inlet / outlet, the chip is considered to be successfully connected to the connector. Therefore, if ΔP_PT>ΔP_P at the inlet (ΔP_PT>ΔP_T at the outlet), the connection can be achieved. In addition Furthermore, if r remains constant and Z increases, the leakage pressure is expected to be smaller. Therefore, to create a larger difference between ΔP_PT and ΔP, ΔP_PT needs to be maximized by decreasing Z and ΔP needs to be minimized by increasing r. This can be achieved by 2r[1+(cosθ_P) / (cosθ_T)]>Z. It should be noted that after successful connection, the robustness of the connection under different inlet and outlet radii was further evaluated in the evaluation of the examples described below. If at the entrance r_P > r_T (or at the exit r_P < r_T), the fluid flowing out of the upstream channel is first pinned at the edge of the Teflon coating (or the surface of the PMMA substrate), and then grows until it reaches the inner wall of the downstream channel (Figs. 2ci and 2di). Although the Teflon coating has a certain thickness (∼20 μm), it is still too thin compared to the droplet size, so the growing fluid will directly bypass the coating. Due to the hydrophilicity of the PMMA substrate (or Al), the fluid will wet the inner wall of the downstream channel and then reach the edge of the downstream channel. Due to the design of the entrance and exit edges (open angle, α is 90°), the fluid will be pinned in the gap again. Since the fluid is pinned by both the surface of the Teflon coating and the edge of the PMMA substrate (or the surface of the Teflon coating and the inner wall and the surface of the PMMA substrate), a larger Z is allowed. When the radius of the upstream channel is smaller than that of the downstream, the maximum leakage pressure is higher than the opposite case (entrance: 2633 KPa vs 1282 KPa, exit: 9578 KPa vs 3298 KPa). In addition, due to the hydrophobicity and roughness of the Teflon coating, the fluid at the exit needs to overcome the local Young-Laplace pressure of the coating inner wall. Therefore, the leakage pressure at the exit is significantly higher than that at the entrance (r_T ∼ 960 μm, exit: 9578 KPa > entrance: 2633 KPa) (Fig. 5a). Conversely, if at the entrance r_P < r_T (or at the exit r_P > r_T), the fluid will grow until it reaches the edge of the downstream channel rather than the inner channel wall. Therefore, the fluid will first wet the PMMA (or Teflon) surface and then move outward until it reaches the maximum contact angle (Figs. 2cii and 2dii). As mentioned above, to prevent leakage, ΔP_PT needs to be higher than ΔP downstream, which makes the acceptable range of Z narrower. The same as the previous example, the fluid has difficulty passing through the surface of the Teflon coating, resulting in the maximum leakage pressure at the exit (3298 KPa) being significantly higher than that at the entrance (1282 KPa) (Fig. 5b). From the above viewpoints, in the biomaterial-plastic composite device, the distance of the gap between the microfluidic chip and the connector is 60 - 335 μm, preferably 60 μm. The thickness of the surface-rough Teflon coating is 20 μm or less. Examples The following examples are shown to further illustrate the present application. <Manufacture of hydrogel-plastic composite chip> A physically cross-linked hydrogel was prepared by dissolving 2.3 g of acrylamide, 5.1 mg of N,N'-methylenebisacrylamide, 20 mg of Irgacure 2959, and 0.15 g of agar in 10 ml of deionized water. The mixture was heated to boiling and stirred until completely dissolved. The gel solution was poured into a sterilized petri dish and allowed to solidify. The solidified hydrogel was cut into the optimal size for bonding. The plastic substrate was prepared by numerical control or soft lithography. For soft lithography, the plastic substrate was placed on a male metal mold in a hot press. The substrate was slightly pressed and formed at a temperature above its melting point and then cooled to room temperature. It should be noted that pressure must be maintained throughout the thermoforming process. The plastic substrate was cleaned with ethanol and water, then immersed in a 10% w / v benzophenone-ethanol solution for 10 minutes, and the substrate was dried with nitrogen. The physically cross-linked hydrogel was placed on top of it. After exposure to ultraviolet light for 5 minutes, the hydrogel underwent chemical cross-linking to form a polyacrylamide hydrogel and was bonded to PMMA simultaneously. Finally, the hydrogel-plastic composite chip was immersed in deionized water to remove all unreacted reactants. <Teflon-coated chip connector> The bottom of the connector was sprayed with Teflon primer and heated at 150 °C for 15 minutes. Then, the Teflon topcoat was sprayed on the connector and sintered at 350 °C for 15 minutes. The surface characteristics of the coating were characterized by scanning electron microscopy. <Connection robustness test> Through holes with different diameters were drilled in the PMMA substrate and the aluminum material with Teflon coating, respectively. Then, they were aligned under a microscope and pressed together for connection. Polytetrafluoroethylene gaskets with different thicknesses were pressed between them to form gaps with different distances. The inlet of PMMA (or Teflon) was connected to a Teflon tube and a syringe barrel, and the outlet of Teflon (or PMMA) was blocked with tape. The syringe barrel was slowly lifted until leakage occurred. The leakage height was measured and used to calculate the leakage pressure. The diameter of the drilled hole and the gap distance were measured under a microscope. The initial gap distance (without polytetrafluoroethylene gasket) was measured using rhodamine B under a confocal microscope. The pore diameters of the aluminum material with Teflon coating were 372 μm and 961 μm, respectively. The pore diameter of the PMMA substrate was 594 μm. The measured gaps were 60 μm (initial), 87 μm, 122 μm, 164 μm, 198 μm, 233 μm, 275 μm, 336 μm, 385 μm, 479 μm, and 603 μm. Regarding the specific results, as described in the above specification, and can be referred to Figure 5a and Figure 5b. <Reliability test (Antibiotic susceptibility test (AST))> To further demonstrate the reliability of the contactless interconnection strategy of the present invention, an antibiotic susceptibility test (AST) was performed, thereby demonstrating that the microfluidic chip and connection method of the present invention can be applied to biological applications. First, the antibiotic solution was diluted in MH broth (MHB), a microbial growth medium, and the microfluidic chip was immersed in MHB, resulting in a uniform MHB concentration in the hydrogel and the antibiotic solution. This ensured a uniform and stable MHB concentration throughout the microfluidic chip, preventing variations in bacterial growth during testing. Green fluorescent protein-expressing Escherichia coli (Gram-negative) and Gram-positive Staphylococcus aureus were used as bacterial samples. MHB medium and antibiotic solutions (diluted to appropriate concentrations in 2.1% w / v MHB) were injected into parallel channels for 1D AST. Alternatively, MHB medium and the two antibiotic solutions were injected into adjacent channels for 2D AST. The bacterial sample was diluted to an appropriate density (OD600 = 0.05). After a stable drug diffusion gradient was generated within the microfluidic chip, a drop of sample solution (5-6 μl) was dripped onto the hydrogel portion of the microfluidic chip. A sterile Petri dish lid was used to cover the entire device to isolate it from the surrounding environment. The device was then incubated at 37.5°C. After 3-4 hours, images were captured using a fluorescence microscope (for E. coli) or an optical microscope (for Staphylococcus aureus) and analyzed using ImageJ 1.4 software to determine the minimum inhibitory concentration (MIC) of the antibiotic. Unlike conventional AST microfluidics, the biomaterial-plastic composite device of this application allows bacterial samples to be added directly to the hydrogel surface rather than being perfused into the channel. This cell-top placement strategy avoids the effects of shear flow and channel contamination. Therefore, the biomaterial-plastic composite device of this application can track the morphology of individual cells and mimic in vivo conditions. In addition, the biomaterial-plastic composite device of the present application was tested against three different classes of antibiotics based on data from the World Health Organization, and their MIC values were compared with data from the Clinical and Laboratory Standards Institute (CLSI) of the United States. AST results depend on the type of antibiotic. For ampicillin, a β-lactam antibiotic, the effect is to inhibit bacterial cell wall synthesis, leading to cell fragility and ultimately bacterial lysis. Therefore, when bacteria show abnormal growth, AST results are measured as an indicator of inhibition. For chloramphenicol, an aminophenol antibiotic, it works by blocking the peptidyl transferase enzyme on the bacterial 50S ribosomal subunit. For gentamicin, an aminoglycoside antibiotic, it works by blocking the peptidyl transferase enzyme on the bacterial 30S ribosomal subunit and contains amino-modified glycosides. Therefore, since both chloramphenicol and gentamicin inhibit protein synthesis, AST results are measured before bacterial growth has completely ceased. To calculate the MIC value, image analysis tools are used to locate the point of inhibition; the precise determination method depends on the bacterial species. The results showed that in 1D AST, the MIC values of ampicillin, chloramphenicol, and gentamicin against Escherichia coli were 2.4 μg / ml, 6 μg / ml, and 0.8 μg / ml, respectively, and against Staphylococcus aureus were 1.8 μg / ml, 5 μg / ml, and 0.3 μg / ml, respectively. These MIC values are consistent with CLSI data. In 2D AST, ampicillin and gentamicin, which have shown synergistic effects in clinical applications, were selected as test drugs. Drug solutions were added to adjacent channels, and Escherichia coli expressing green fluorescent protein was used as a representative bacterial sample. When a stable concentration gradient of the two drugs was established, the drugs began to affect bacterial growth. Using this device, users can observe single-cell morphological changes caused by the drug action mechanism, as shown in Figure 6a. When cells approached the area with high gentamicin concentration and low ampicillin concentration, cell division ceased, and the cell number remained unchanged (Figure 6b). When cells were in the area with medium gentamicin concentration and low ampicillin concentration, the drug concentration failed to inhibit cell division, causing the cell number to begin to increase (Figure 6c). When cells are in the medium gentamicin concentration range and the high ampicillin concentration range, ampicillin inhibits bacterial cell wall synthesis, causing cells to elongate and rupture, while gentamicin maintains cell number (Figure 6d). When cells are in the low concentration range of both drugs, cell morphology is unaffected, and cells divide normally, increasing in number (Figure 6e). When cells are in the low ampicillin concentration range and the gentamicin concentration is slightly below the MIC, the gentamicin concentration is insufficient to completely halt cell division, but it does affect the division rate, causing a slight decrease in cell number (Figure 6f). When cells are in the medium concentration range of both drugs, ampicillin causes cells to begin to elongate and impairs cell division, leading to a decrease in cell number (Figure 6g). However, this elongation can easily create the visual illusion of an increase in cell number, so it is important to also pay attention to changes in cell number. Finally, when cells are in the low gentamicin concentration range and the high ampicillin concentration range, cell elongation and an increase in number occur (Figure 6h). This is due to the low gentamicin concentration, which is insufficient to produce a synergistic effect. Therefore, in cases where only a single drug is effective (as in 1D AST), a higher concentration of ampicillin is required to affect cell proliferation. Compared to traditional methods, the biomaterial-plastic composite device of the present application can obtain MIC values within 3-4 hours, significantly accelerating diagnosis. Furthermore, the biomaterial-plastic composite device of the present application demonstrates potential for clinical applications, including but not limited to rapid AST. As mentioned above, regarding concentration, expressions such as "low", "medium", "high", and "slightly below the MIC value" are used. It should be noted that the concentration relationship represented by these expressions is "low" < "medium" < "slightly below the MIC value" < "high". As shown in the AST test results, the composite material chip design not only enhances the mechanical strength of the hydrogel chip, but also enjoys the advantage of cell top placement, ensuring the authenticity of the test results. Industrial Applicability The design of the microfluidic chip of the present application not only enhances the mechanical strength of the hydrogel chip, but also enjoys the advantage of cell top placement, ensuring the authenticity of the test results. The microfluidic chip of the present application can still be connected to the connector using the pinning effect even in the presence of a gap, thereby realizing a contactless interconnection strategy. The contactless interconnection method between the microfluidic chip of the present application and the connector is very simple, does not require precise manufacturing process or alignment operation (allowing connection with gaps), and can be mass-produced. Therefore, the bottleneck of hydrogel chip connection is overcome, automation and large-scale production are realized, and the microfluidic chip is more suitable for clinical use.

Claims

1. A microfluidic chip, which is a composite material chip obtained by arranging biomaterials on a plastic substrate, characterized in that: A microfluidic channel is constructed on the plastic substrate.

2. The microfluidic chip according to claim 1, wherein The biomaterial is polyacrylamide hydrogel, or a combination of polyacrylamide hydrogel and natural materials or peptides.

3. The microfluidic chip according to claim 2, wherein: The biomaterial is polyacrylamide hydrogel.

4. The microfluidic chip according to claim 2 or 3, wherein: The natural material is agar or agarose or alginate or collagen.

5. The microfluidic chip according to claim 1, wherein The plastic substrate is formed of polymethyl methacrylate, polystyrene, polypropylene, or polyethylene.

6. The microfluidic chip according to claim 5, characterized in that The plastic substrate is formed of polymethyl methacrylate.

7. A biomaterial-plastic composite material device, comprising the microfluidic chip and the connector according to any one of claims 1 to 6, characterized in that: The microfluidic channel on the plastic substrate has ports serving as an inlet and an outlet, respectively. The channel on the connector has ports serving as an inlet and an outlet, The fluid flows in the biomaterial-plastic composite material device in the order of the outlet of the connector, the inlet of the microfluidic channel, the outlet of the microfluidic channel, and the inlet of the connector. The surface of the connector is formed with a Teflon coating, There is a gap between the microfluidic chip and the connector, that is, the connection between the two is non-contact.

8. The biomaterial-plastic composite material device according to claim 7, wherein: The distance of the gap is 60 to 335 μm.

9. The biomaterial-plastic composite material device according to claim 8, wherein: The distance of the gap is 60 μm.

10. The biomaterial-plastic composite material device according to claim 7, wherein: The radius of the outlet of the connector is smaller than the radius of the inlet of the microfluidic channel, The radius of the inlet of the connector is larger than the radius of the outlet of the microfluidic channel.

11. A non-contact interconnection method between a microfluidic chip and a connector, characterized in that: The microfluidic chip is the microfluidic chip according to any one of claims 1 to 6, The microfluidic channel on the plastic substrate of the microfluidic chip has ports serving as an inlet and an outlet, respectively. The channel on the connector has ports serving as an inlet and an outlet, The fluid flows in the biomaterial-plastic composite material device in the order of the outlet of the connector, the inlet of the microfluidic channel, the outlet of the microfluidic channel, and the inlet of the connector. The surface of the connector is formed with a Teflon coating, There is a gap between the microfluidic chip and the connector.

Images