Microfluidic device and method for introducing microparticles into cells

Abstract

The application discloses a microfluidic device and a method for introducing microparticles into cells. The method comprises injecting a microparticle solution and a cell suspension into an inlet channel, and the microparticle solution and the cell suspension are mixed and flow into a main channel. When flowing through the main channel, gas in a microstructure and mixed liquid in the main channel form a gas-liquid interface at a first gap. A piezoelectric transducer is controlled to be turned on, and the piezoelectric transducer generates sound waves. The sound waves make the gas-liquid interface vibrate and generate acoustic vortices near the gas-liquid interface. The acoustic vortices drive the cells and the microparticles in the mixed liquid to gather and rotate. The cells reversibly deform under stress to generate micropores on the cell membrane. The microparticles enter the cells through the micropores on the cell membrane, and the process of loading the cells with the microparticles is completed. The device and the method disclosed by the application realize higher microparticle loading efficiency and solve the problem of uneven microparticle loading effect while protecting the integrity and functionality of the cells.

Description

A microfluidic device and a method for introducing microparticles into cells Technical Field

[0001] This invention relates to the field of biomedicine, specifically to a microfluidic device and a method for introducing microparticles into cells. Background Technology

[0002] Cellular drug delivery systems utilize the unique physiological functions of cells to target drugs to diseased areas, overcoming the limitations of traditional drug delivery systems and showing promise as a new generation of intelligent drug delivery systems. Current methods to enhance cell drug loading include co-incubation, electroporation, physical compression, or ultrasound; however, these methods suffer from low loading efficiency and are prone to damaging cell integrity and function.

[0003] Chinese patent CN110560186A discloses a method for synthesizing biomembrane nanoparticles using a microfluidic chip, and the microfluidic chip itself. The method includes: introducing a biomembrane solution into a biomembrane channel; introducing an organic solution containing a polymer into a polymer channel to mix the organic solution and the biomembrane solution; and introducing a PBS solution into a PBS channel to mix the PBS solution with the mixed solution. The mixed solution is then output to a second mixing channel, whereby the polymer in the mixed solution is forced into the biomembrane under pressure and ultrasound. This invention combines a microfluidic chip with ultrasound to provide a one-step method for synthesizing biomembrane-encapsulated nanoparticles. By using a microfluidic chip, the flow rate and concentration of materials can be precisely controlled to regulate the size and concentration of the nanoparticles, resulting in synthesized biomembrane nanoparticles with uniform morphology and stable structure.

[0004] While the aforementioned patented technologies have successfully achieved the regulation of nanoparticle size and concentration, as well as the uniform synthesis and stable structure of biomembrane nanoparticles through precise control of material flow rate and concentration, challenges remain in practical applications. Specifically, the small channel size of microfluidic chips results in extremely low Reynolds numbers for the fluid within them. Under low Reynolds number conditions, the fluid exhibits a stable stratified flow state, where viscous forces are much greater than inertial forces, and the flow exhibits a parallel streamline distribution with a lack of macroscopic turbulent disturbances. This flow characteristic limits the contact between particles and cells, allowing them to contact only through diffusion. However, the diffusion distance is limited by the micrometer-scale channel size, resulting in insufficient mixing of particles (such as drug nanoparticles) and cells within the microchannels, leading to significant fluctuations in drug loading efficiency. Summary of the Invention

[0005] The purpose of this invention is to provide a microfluidic device and a method for introducing microparticles into cells, so as to solve the technical problems mentioned in the background art.

[0006] To achieve the above objectives, the present invention provides the following technical solution:

[0007] The first aspect of the present invention provides a microfluidic device, which includes a microfluidic chip and a piezoelectric transducer. The microfluidic chip includes a cavity channel, which includes an inlet channel, a main channel, and an outlet channel. At least one microstructure is formed on the sidewall of the main channel, and the microstructure is connected to the main channel through a first gap. The microstructure is used to form a gas-liquid interface with the liquid in the main channel at the first gap when the liquid flows through the main channel. The piezoelectric transducer is used to generate sound waves, which cause the gas-liquid interface between the liquid in the main channel and the gas in the microstructure to vibrate and generate acoustic eddies near the gas-liquid interface.

[0008] A second aspect of the present invention provides a method for introducing microparticles into cells, the method employing the microfluidic device provided by the present invention. The method includes: injecting a microparticle solution and a cell suspension into an inlet channel, the microparticle solution and cell suspension converging and mixing to flow into a main channel; as the mixed liquid flows through the main channel, a gas-liquid interface is formed between gas within at least one microstructure and the mixed liquid within the main channel at a first gap; controlling the activation of a piezoelectric transducer, the piezoelectric transducer generating acoustic waves, the acoustic waves causing vibration of the gas-liquid interface between the mixed liquid within the main channel and the gas within the microstructure, and generating acoustic eddies near the gas-liquid interface; the acoustic eddies cause cells and microparticles within the mixed liquid to aggregate and rotate, the cells undergo reversible deformation under force, generating micropores on the cell membrane, and the microparticles enter the cell through the micropores on the cell membrane.

[0009] The technical effects and advantages of this invention are as follows:

[0010] The microfluidic device and method for introducing microparticles into cells provided by this invention fully utilize the gas-liquid interface formed within the microfluidic chip to create enhanced acoustic vortices driven by acoustic waves. This allows microparticles and cells flowing through the acoustic vortices to achieve efficient and controllable introduction of microparticles into cells. Simultaneously, the acoustic vortices effectively avoid problems such as cell breakage or uneven microparticle introduction during the process. Furthermore, the microfluidic device-based method offers advantages such as small sample capacity and controllable flow rate. Attached Figure Description

[0011] Figures 1A-B are a top view and a side view of a microfluidic device disclosed in this invention;

[0012] Figures 2A-F are schematic diagrams of the cavity channel and various microstructures disclosed in this invention;

[0013] Figures 3A-F are enlarged views of parts of Figures 2A-F;

[0014] Figures 4A-F are schematic diagrams of the gas-liquid interface formed between the main channel and the various microstructures when the liquid flows through the main channel disclosed in this invention.

[0015] Figures 5A-F are enlarged views of parts of Figures 4A-F;

[0016] Figures 6A-B are a top view and a side view of the second type of microfluidic device disclosed in this invention;

[0017] Figures 7A-B are top and side views of the third microfluidic device disclosed in this invention;

[0018] Figures 8A-B are top and side views of the fourth microfluidic device disclosed in this invention;

[0019] Figures 9A-B are a top view and a side view of the fifth type of microfluidic device disclosed in this invention;

[0020] Figure 10 is a flowchart of a method for introducing microparticles into cells disclosed in this invention;

[0021] Figure 11 shows the experimental results of using the microfluidic device disclosed in this invention to achieve cell drug delivery.

[0022] Figures 12A-B show the application of the flow guide of the present invention when the microstructure is in an asymmetrical arrangement;

[0023] Figures 13A-B show the application of the flow guide of the present invention when the microstructure is symmetrically arranged.

[0024] In the picture:

[0025] 10. Microfluidic devices;

[0026] 1. Microfluidic chip; 101. Chip cover; 102. Chip substrate;

[0027] 2. Cavity passage; 201. Inlet passage; 202. Main passage; 203. Outlet passage;

[0028] 3. Microstructure;

[0029] 4. First gap;

[0030] 5. Piezoelectric transducer;

[0031] 6. Airflow guide section. Detailed Implementation

[0032] The technical solutions of the present invention will be clearly and completely described below with reference to the accompanying drawings of the embodiments of the present invention. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. All other embodiments obtained by those skilled in the art based on the embodiments of the present invention without creative effort are within the scope of protection of the present invention.

[0033] Example 1

[0034] Referring to Figures 1 to 11, the present invention provides a microfluidic device that can improve drug loading efficiency. The microfluidic device 10 includes a microfluidic chip 1, which includes a cavity channel 2. The cavity channel 2 includes an inlet channel 201, a main channel 202, and an outlet channel 203. At least one microstructure 3 is formed on the side wall of the main channel 202, and the microstructure 3 is connected to the main channel 202 through a first gap 4. The microfluidic device 10 also includes a piezoelectric transducer 5, which is mainly used to generate sound waves.

[0035] When the liquid flows through the main channel 202, the gas in the microstructure 3 will form a gas-liquid interface with the liquid in the main channel 202 at the first gap 4, which controls the piezoelectric transducer 5 to work and generate sound waves. The sound waves drive the liquid and gas to vibrate at the gas-liquid interface, thereby generating acoustic eddies near the gas-liquid interface.

[0036] Figure 1 shows a top view and a side view of the microfluidic device 10. It should be noted that the microfluidic device 10 shown in Figure 1 is only a schematic diagram, and the scope of protection of this invention is not limited to the connection relationship between different structures, the number of different structures, the size, etc. in the figure.

[0037] Specifically, the inlet channel 201 may include one, meaning that solutions of different components are injected into the same inlet channel 201. The inlet channel 201 may also include one or more channels, and solutions of different components may be injected into different inlet channels 201 respectively. This invention is not limited herein.

[0038] Specifically, the sidewall of the main channel 202 may be provided with N microstructures 3, where N can be 1 (i.e., only one microstructure 3 is provided on one side) or multiple microstructures 3. The N microstructures 3 may be symmetrically distributed on both sides of the main channel 202 or asymmetrically distributed. Different microstructures 3 may be arranged at equal intervals or at unequal intervals.

[0039] The microstructure 3 can have various shapes. Figures 2A-F show schematic diagrams of various microstructure 3 designs, and Figures 2A-F also show enlarged views of specific parts. Specifically:

[0040] Referring to Figure 2A, the microstructure 3 can adopt a near-circular design, which is equivalent to the sidewall of the main channel 202 intersecting with a circle, and the two points of intersection forming a first gap 4. The microstructure 3 and the main channel 202 are connected through the first gap 4.

[0041] Referring to Figure 2B, the microstructure 3 can adopt a fan-shaped design, which is equivalent to the side wall of the main channel 202 intersecting with the fan shape, and the two points of intersection forming the first gap 4.

[0042] Referring to Figure 2C, another type of fan-shaped structure is shown. The radius and degree of intersection of the fan-shaped structures can be different from the design shown in Figure 2B, thus forming microstructures 3 of different shapes. The length of the first gap 4 formed can also be different.

[0043] Referring to Figure 2D, a rectangular design is adopted, which means that the side wall of the main channel 202 intersects the rectangle at a certain angle, and the two points of intersection form the first gap 4. The angle and length of the intersection can be adaptively adjusted. Figure 2D only shows one example.

[0044] Figure 2E adopts a fan-shaped design, but unlike Figure 2B or Figure 2C, the microstructure 3 is not completely outside the side wall of the main channel 202; the microstructure 3 can also be partially embedded inside the main channel 202.

[0045] Figure 2F also adopts a near-circular design. Unlike Figure 2A, the diameter of the circle and the degree of intersection with the side wall of the main channel 202 can be different, so the first gap 4 formed can be different.

[0046] It should be stated that Figures 2A-F only provide schematic diagrams of several structural designs for the microstructure 3 and do not constitute a limitation on the scope of protection of this invention. It is understood that the microstructure 3 disclosed in this invention can be designed based on basic geometric shapes such as quasi-circular, quasi-sectoral, quasi-rectangular, quasi-trapezoidal, and quasi-rhomboid, and formed through methods such as intersection, rotation, tangency, and tessellation. The first gap 4 corresponding to different microstructures 3 may be different, but it can also be the same. The equivalent areas between microstructures 3 with different structures may be different, but they can also be the same. For N microstructures 3, microstructures 3 with the same structure can be used, and the length of the first gap 4 corresponding to each microstructure 3 is the same. N microstructures 3 can also be designed with different structures, and the length of the first gap 4 between microstructures 3 with different structures is different.

[0047] Figures 4A-F show schematic diagrams of the gas-liquid interfaces formed between the main channel 202 and the N microstructures 3 when liquid flows through the main channel 202 under different microstructure designs 3. Figures 5A-F are enlarged views. It can be seen that when liquid flows through the main channel 202, liquid does not flow into M of the N microstructures 3. The gas in these M microstructures 3 and the liquid in the main channel 202 will form gas-liquid interfaces at the first gap 4, where M is less than or equal to N. This means that it is possible that no liquid flows into any of the N microstructures 3, resulting in N gas-liquid interfaces. Alternatively, liquid may flow into some of the microstructures 3, in which case M gas-liquid interfaces will be formed at the first gap 4 corresponding to the M microstructures 3 (less than N). When the piezoelectric transducer 5 is controlled to start working and generate sound waves, acoustic eddies will be generated near some or all of the M gas-liquid interfaces.

[0048] As shown in Figure 1, by way of example, the microfluidic chip 1 also includes a chip cover 101 and a chip substrate 102. The chip cover 101 may be made of PDMS material, and a cavity channel 2 may be provided in the chip cover 101.

[0049] Figures 1 and 6-9 illustrate examples of different positional relationships between the piezoelectric transducer 5 and the microfluidic chip 1. As shown in Figure 1, the piezoelectric transducer 5 can be installed at the bottom of the chip substrate 102; as shown in Figure 6, the piezoelectric transducer 5 can be installed on the upper part of the chip cover 101; as shown in Figure 7, the piezoelectric transducer 5 can be connected to the upper side of the chip substrate 102 and located on one side of the main channel 202 sidewall; as shown in Figure 8, the piezoelectric transducer 5 can be connected to the upper side of the chip substrate 102 and close to the outlet channel 203; as shown in Figure 9, the piezoelectric transducer 5 can also be connected to the upper side of the chip substrate 102 and close to the inlet channel 201.

[0050] The piezoelectric transducer 5 is used to generate sound waves, which in turn excite the formation of acoustic vortices at the gas-liquid interface between the microstructure 3 and the main channel 202. The piezoelectric transducer 5 and the microfluidic chip 1 are positioned at different locations to control the generation of acoustic vortices with different effects.

[0051] Figure 10 shows a flowchart of a method for introducing microparticles into cells according to the present invention. The method can be implemented using the microfluidic device 10 disclosed in the present invention. The method is described below:

[0052] S1. Inject microparticle solution and cell suspension into inlet channel 201. The microparticle solution and cell suspension converge and mix before flowing into main channel 202.

[0053] It should be noted that the inlet channel 201 can be designed as a single channel, in which case the microparticle solution and cell suspension are injected into the same inlet channel 201 and converge and mix as they flow through the inlet channel 201 and the main channel 202. The resulting mixed liquid then flows through the main channel 202. Alternatively, the inlet channel 201 can be designed as two channels, in which case the microparticle solution and cell suspension are injected into different inlet channels 201 and converge and mix at the junction of the different inlet channels 201. The resulting mixed liquid then flows through the main channel 202. It is understood that the number of inlet channels 201 can be designed according to the types of solutions to be mixed, and this invention does not impose any limitation on this.

[0054] Microparticle solutions can be sample solutions containing drugs, gene fragments, or plasmids of appropriate size. Cell suspensions can be PBS solutions containing cells such as red blood cells, immune cells, stem cells, or tumor cells.

[0055] S2. When the mixed liquid flows through the main channel 202, the gas in at least one microstructure 3 and the mixed liquid in the main channel 202 form a gas-liquid interface at the first gap 4.

[0056] If the outer wall of the main channel 202 is provided with N microstructures 3, when the mixed liquid flows through the main channel 202, there will be at least M microstructures 3 in which no mixed liquid flows in, where M is less than or equal to N. At this time, gas is present in all M microstructures 3, and a gas-liquid interface will be formed at the first gap 4 between these M microstructures 3 and the main channel 202.

[0057] S3. Control the piezoelectric transducer 5 to turn on. The piezoelectric transducer 5 generates sound waves. The sound waves cause the gas-liquid interface between the mixed liquid in the main channel 202 and the gas in the microstructure 3 to vibrate, generating acoustic eddies near the gas-liquid interface.

[0058] S4. Acoustic eddies cause cells and particles in the mixed liquid to gather and rotate. The cells undergo reversible deformation under force, creating micropores on the cell membrane. Particles enter the cell through the micropores on the cell membrane.

[0059] S5. After the preset loading time, control the piezoelectric transducer 5 to close, and the mixed liquid continues to flow in the main channel 202 until it flows out of the outlet channel 203.

[0060] The microfluidic device 10 includes a piezoelectric transducer 5. By controlling the opening and closing of the piezoelectric transducer 5, the generation of acoustic vortices and the continuous flow of the mixed liquid within the main channel 202 can be controlled. When the piezoelectric transducer 5 is turned on, a signal generator produces a sine wave of a specific frequency, which, after being amplified by a signal amplifier, drives the transducer to generate sound waves. The sound wave range includes, but is not limited to, 1kHz-1GHz. Since the piezoelectric transducer 5 is partially connected to the microfluidic chip 1, the sound waves can be transmitted to the cavity channel 2 of the microfluidic chip 1. Under the excitation of the sound waves, the gas-liquid interface formed between the mixed liquid in the main channel 202 and the gas in the microstructure 3 will vibrate, and acoustic vortices will be generated near the gas-liquid interface. Figure 11 shows the experimental results of using the microfluidic device 10 disclosed in this invention to achieve cell drug delivery. As can be seen from Figure 11, different acoustic vortices are formed at the gas-liquid interface between different microstructures 3 and the main channel 202.

[0061] The generated acoustic eddies cause cells and particles in the liquid mixture to aggregate and rotate. Under the influence of fluid drag and centrifugal force, the cells undergo reversible deformation, creating micropores on their cell membranes. Particles in the liquid mixture, such as drugs, gene fragments, or plasmids, can enter the cells through these micropores, achieving the process of introducing particles into the cells, such as drug delivery. During this process, an electric field or other energy source can be applied to further stimulate and enlarge the surface micropores, thereby enhancing particle introduction efficiency or allowing larger particles to enter the cells.

[0062] After the preset particle loading time has elapsed, the piezoelectric transducer 5 is turned off. At this time, the mixed liquid in the main channel 202 continues to flow along the channel direction until it flows out of the outlet channel 203, where the sample is collected, completing the cell loading particle process. The set time period can be from milliseconds to seconds and can be adaptively adjusted according to the particle loading effect; this invention does not limit the time period.

[0063] In the microfluidic chip 1 disclosed in this invention, multiple microstructures 3 are provided on the sidewall of the main channel 202, which can form multiple sets of acoustic vortex excitation units and generate multiple sets of acoustic vortices. Different flow rates and particle loading times can be designed as needed, so that the sample can undergo multiple rounds of repeated loading processes while flowing through multiple sets of acoustic vortices, thereby improving loading efficiency, loading effect and stability.

[0064] When the piezoelectric transducer 5 is turned off, the mixed liquid flows along the channel direction in the main channel 202 within a preset off time. After the preset off time, the piezoelectric transducer 5 can be turned on again. Under the action of the acoustic waves generated by the piezoelectric transducer 5, acoustic vortices are generated again at the multiple gas-liquid interfaces corresponding to the multiple microstructures 3. Driven by the acoustic vortices, the cells and particles in the mixed liquid gather and rotate again. The cells undergo reversible deformation under force, generating micropores on the cell membrane. The particles enter the cells through the micropores of the cell membrane. After the preset loading time, the piezoelectric transducer 5 is turned off, and the mixed liquid continues to flow along the channel direction in the main channel 202. It can be understood that during the process of the sample flowing through the entire main channel 202 from entering the inlet channel 201 to flowing out of the outlet channel 203, the cell and particle loading process can be realized once or multiple times by controlling the opening or closing of the piezoelectric transducer 5, until the sample flows out of the outlet channel 203 and the sample collection is completed, thus completing the entire process of cell and particle loading. Different flow rates, number of cell particle loading times, and loading times can be designed as needed, and this invention does not limit these parameters.

[0065] This invention discloses a microfluidic device 10 and a method for introducing microparticles into cells based on the designed and fabricated microfluidic device 10. By using acoustic waves to drive the vibration of the designed gas-liquid interface in the microfluidic chip 1, a regular and enhanced acoustic vortex is formed. The designed microfluidic chip 1 guides the microparticles and cells containing the microparticles to flow through the acoustic vortex. Under the action of the acoustic vortex, a highly efficient and controllable cell loading process of microparticles is achieved, effectively improving the efficiency of microparticle loading into cells and enabling precise control of the microparticle loading amount. The advantage of this invention lies in controlling the generation, frequency, and intensity of the enhanced acoustic vortex using acoustic waves, thereby controlling the occurrence of the cell loading process of microparticles. Compared with other methods of enhancing drug loading effects such as electroporation, physical extrusion, or ultrasound, it has better controllability, higher efficiency (i.e., higher drug loading efficiency), and more precise loading amount control. Simultaneously, the application of microfluidic technology also allows for better control of the sample flow rate and volume, enabling better control of the consistency of the loading amount per cell during mass production. In the field of cell drug delivery, it solves the problems and shortcomings of traditional methods, such as long loading time, low efficiency, poor precision, poor consistency and difficulty in controlling drug loading, while achieving higher drug loading efficiency than traditional methods.

[0066] Example 2

[0067] Although the above embodiments can improve drug loading efficiency through the combination of microfluidic device and piezoelectric transducer 5, in practical applications, the small channel size of microfluidic chip 1 results in extremely low Reynolds number of the fluid. Under low Reynolds number conditions, the fluid exhibits a stable stratified flow state, where the viscous force is much greater than the inertial force, the flow is distributed with parallel streamlines, and there is a lack of macroscopic turbulent disturbance. This flow characteristic limits the contact mode between particles and cells, making them only able to contact each other through diffusion. Although drug loading efficiency can be improved through acoustic eddies, the generation of acoustic eddies in the above embodiments is affected by the working state of microstructure 3 and cannot fully cope with cell-loaded particles. At the same time, the influence range of acoustic eddies is small, resulting in particles and cells located in the middle of the main channel not being able to mix by acoustic eddies. This leads to an improvement in drug loading efficiency, but it is still not comprehensive. In view of this, a technical improvement is made based on embodiment one. The improved technical solution is as follows:

[0068] Referring to Figures 1 to 13, the present invention provides a microfluidic device. The microfluidic device 10 includes a microfluidic chip 1 and a first gap 4. The microfluidic chip 1 has a cavity channel 2 inside. The cavity channel 2 includes an inlet channel 201, a main channel 202 and an outlet channel 203. The mixed liquid enters through the inlet channel 201, completes the mixing of microparticles and cells in the main channel 202, and then exits through the outlet channel 203.

[0069] The main channel 202 has microstructures 3 on its side wall. The microstructures 3 are arrayed on the side wall of the main channel 202. The microstructures 3 located on both sides of the main channel 202 can be symmetrically or staggeredly distributed.

[0070] The microstructure 3 is connected to the main channel 202 through the first gap 4. The straight line formed by the microstructure 3 and the first gap 4 is perpendicular to or intersects the main channel 202. When the straight line formed by the microstructure 3 and the first gap 4 is perpendicular to the main channel 202, the size of the first gap 4 is smaller than the size of the microstructure 3. When the liquid enters the main channel 202, due to the small size of the first gap 4, it is difficult for the gas inside the microstructure 3 to escape. This makes it easier for a gas-liquid interface to form between the first gap 4 of the microstructure 3 and the main channel 202 during the liquid flow.

[0071] When the straight line formed by the microstructure 3 and the first gap 4 intersects with the main channel 202, the first gap 4 faces the outlet channel 203. When the liquid enters the main channel 202, due to the influence of the liquid flow direction and the orientation of the first gap 4, the liquid is difficult to penetrate into the microstructure 3, thus making it easier for a gas-liquid phase interface to form between the first gap 4 of the microstructure 3 and the main channel 202 during the liquid flow.

[0072] When the liquid flows through the main channel 202, the gas in the microstructure 3 and the liquid in the main channel 202 form a gas-liquid interface at the first gap 4; the sound waves generated by the piezoelectric transducer 5 when it is working generate acoustic eddies near the gas-liquid interface, which in turn cause the cells and particles in the liquid to gather and rotate. The cells undergo reversible deformation under force, generating micropores on the cell membrane, and the particles enter the cell through the micropores of the cell membrane.

[0073] A flow guide 6 is provided on the outside of the first gap 4. The flow guide 6 is used to guide the liquid to the first gap 4. The length of the flow guide 6 located on the outside of the first gap 4 is less than or equal to the length between the two first gaps 4. There is an angle X between the flow guide 6 and the main channel 202. When X is 0, the flow guide 6 coincides with the side wall of the main channel 202, and the side wall of the main channel 202 guides the direction of liquid flow.

[0074] When X is greater than 0, the guide section 6 forms a relatively inclined state on both sides of the first gap 4. Specifically, the guide section 6 on the side of the first gap 4 facing the inlet channel 201 is deflected towards the middle of the main channel 202; the guide section 6 on the side of the first gap 4 facing the outlet channel 203 is deflected towards the side wall of the main channel 202. That is, when X is greater than 0, the guide section 6 is used to guide the liquid to the other side of the first gap 4.

[0075] Take Figure 12A-B as an example:

[0076] When the microstructures 3 on both sides of the main channel 202 are staggered, the length of the guide portion 6 located outside the first gap 4 is less than the length between two adjacent first gaps 4 on the same side wall of the main channel 202. The guide portion 6 forms a serpentine distribution in the main channel 202, as shown in Figure 12B.

[0077] When in use, the sound waves generated by the piezoelectric transducer 5 drive the gas and liquid to generate an acoustic vortex at the first gap 4. After the liquid flows into the acoustic vortex, the acoustic vortex causes the cells and particles in the mixed liquid to gather and rotate. The cells undergo reversible deformation under force, generating micropores on the cell membrane. The particles enter the cell through the micropores on the cell membrane.

[0078] After the liquid exits the acoustic vortex, it is affected by viscous resistance and gradually returns to a laminar flow state. Under the driving force of the subsequent liquid, the liquid continues to flow towards the outlet channel 203. At the same time, since the guide part 6 forms a relatively inclined state on both sides of the first gap 4, the guide part 6 guides the liquid to the other side of the first gap 4. As a result, when the liquid flows to the guide part 6, it will be subjected to a certain deflection force, causing it to move towards the side wall of the main channel 202. As a result, the particles originally located in the middle are also concentrated towards the side wall of the main channel 202 due to the deflection force of the liquid. Since the microstructures 3 on both sides of the main channel 202 are staggered, after the liquid flows to the microstructures 3 on the other side, the particles are more concentrated. Therefore, under the influence of the acoustic vortex at the microstructures 3, the micropores on the cell membrane of the particles are more likely to enter the cell, further improving the drug loading rate of the cells.

[0079] Meanwhile, as shown in Figure 12A, the microstructure 3 is not entirely on the outer side of the main channel 202; rather, it is partially embedded inside the main channel 202, with the flow guides 6 symmetrically distributed on both sides of the microstructure 3. Based on this arrangement, when the liquid flows through the main channel 202, the vertical distance between the first gap 4 and the main channel 202 is less than the width of the main channel 202. Therefore, when the liquid flows to the area between the first gap 4 and the main channel 202, the acoustic eddies generated at the first gap 4 can capture more particles, thus ensuring the possibility of particles entering the cell through the micropores on the cell membrane and improving the drug loading rate of the cell.

[0080] Take Figure 13A-B as an example:

[0081] When the microstructures 3 on both sides of the main channel 202 are symmetrically distributed, the length of the guide portion 6 located outside the first gap 4 is equal to the length between two adjacent first gaps 4 on the same side wall of the main channel 202. The guide portion 6 forms a contraction-diffusion arrangement in the main channel 202, as shown in Figure 13B.

[0082] During use, after the liquid enters the main channel 202, it first undergoes the first cell drug loading through the acoustic vortex at the first symmetrically distributed gap 4. Then, influenced by the constricting guide section 6, the liquid converges towards the center. During this process, the guide section 6, influenced by the direction of liquid flow, gathers the particles distributed in the side wall area of ​​the main channel 202 towards the center, thereby reducing the distance between cells and particles, facilitating full contact between cells and particles, and preventing particles from being too dispersed in the liquid, which would prevent cells from fully contacting particles and thus reduce the drug loading capacity of the cells.

[0083] When the liquid flows to the boundary between the contraction zone and the diffusion zone, the acoustic vortex at the first gap 4 on both sides of the main channel 202 at this location performs a second cell drug loading. During this process, because the particles are relatively concentrated, the acoustic vortex generated at the first gap 4 can capture more particles. At the same time, because the distance between the two first gaps 4 is relatively close, the outer regions of the acoustic vortex at both locations can come into contact with each other. This allows the cells located in the outer region of the acoustic vortex to also capture a considerable number of particles. This avoids the phenomenon that the central region of the acoustic vortex exerts a strong force on the cells, making it easy for them to capture particles, while the outer region of the acoustic vortex exerts a weak force on the cells, making it difficult for the cells distributed in the outer region to undergo reversible deformation to generate micropores, thus reducing the possibility of particles entering the cells through the cell micropores.

[0084] When fluid enters the diffusion zone from the narrow contraction zone, the increased cross-sectional area causes a sharp drop in velocity. At this point, some kinetic energy is converted into pressure energy, and the local pressure rises, forming a reverse pressure gradient (pressure increases along the flow direction). The fluid near the wall has a lower velocity due to viscosity and cannot resist the pressure increase under the reverse pressure gradient, causing the boundary layer to detach from the wall, forming a low-pressure backflow vortex. This enhances turbulent mixing and disrupts laminar flow stability. Consequently, the particles entering the diffusion zone can fully contact and mix with the cells. When the liquid flows to the boundary between the diffusion and contraction zones, the acoustic vortex generated at the first gap 4 can more easily deliver particles through the cell membrane micropores into the cells, improving the cell's drug loading efficiency. By simultaneously designing the flow channel structure of the diffusion and contraction zones, the problem of particle aggregation caused by continuous contraction flow in the microfluidic channel is effectively solved. When the fluid flows through the contraction section, the particles migrate towards the center of the main channel 202 under the action of inertial force; the introduction of the diffusion zone can interrupt this migration process in time, avoiding the formation of high concentrations of particles in the subsequent narrow areas, especially at geometrically abrupt locations such as bends or bifurcations, significantly reducing the risk of particle blockage caused by local velocity gradient changes, and ensuring the long-term stable operation of the microfluidic system.

[0085] By alternately setting the contraction zone and the diffusion zone, the liquid velocity increases in the contraction zone, and the cells and particles migrate towards the center of the main channel 202 due to inertia; the velocity drops sharply in the expansion zone, generating eddies or secondary flows, which disrupt the laminar flow boundary and promote the radial mixing of cells and particles; in particular, the acoustic eddies generated by the piezoelectric transducer 5 through the first gap 4 enable the drug loading rate of the cells to be comprehensively improved.

[0086] Finally, it should be noted that the above description is only a preferred embodiment of the present invention and is not intended to limit the present invention. Although the present invention has been described in detail with reference to the foregoing embodiments, those skilled in the art can still modify the technical solutions described in the foregoing embodiments or make equivalent substitutions for some of the technical features. Any modifications, equivalent substitutions, improvements, etc., made within the spirit and principles of the present invention should be included within the protection scope of the present invention.

[0087] Although embodiments of the invention have been shown and described, those skilled in the art will recognize that various changes, modifications, substitutions and variations can be made to these embodiments without departing from the principles and spirit of the invention, the scope of which is defined by the appended claims and their equivalents.

Claims

1. A microfluidic device, comprising a microfluidic chip, characterized in that: A cavity channel, located inside the microfluidic chip, includes an inlet channel, a main channel, and an outlet channel; a microstructure, located on the side wall of the main channel and connected to the main channel through a first gap; when liquid flows through the main channel, the gas in the microstructure and the liquid in the main channel form a gas-liquid interface at the first gap; a guide section, located outside the first gap, guides the liquid towards the first gap; a piezoelectric transducer generates sound waves, which enable the gas-liquid interface to vibrate and generate acoustic eddies near the gas-liquid interface, thereby causing cells and particles in the liquid to aggregate and rotate. The cells undergo reversible deformation under force, creating micropores on the cell membrane, through which particles enter the cell. The number of microstructures set on the side wall of the main channel is N, where N is greater than or equal to 1. When liquid flows through the main channel, if no liquid flows into M of the microstructures, where M is less than or equal to N, then the gas in the M microstructures and the liquid in the main channel form gas-liquid interfaces at the M first gaps respectively. When the number of microstructures on both sides of the main channel is greater than 1 and they are symmetrically distributed, the length of the guide part located outside the first gap is equal to the length between two adjacent first gaps. The guide part forms a contraction-diffusion arrangement in the main channel. The microstructures are embedded inside the main channel, and the guide parts are symmetrically distributed on both sides of the microstructures.

2. The microfluidic device according to claim 1, characterized in that, The N microstructures have the same structure, and the length of the first gap corresponding to each microstructure is the same.

3. The microfluidic device according to claim 1, characterized in that, The structures of the N microstructures are not completely identical, and the lengths of the first gaps between the microstructures with different structures are not the same.

4. The microfluidic device according to claim 1, characterized in that, The microfluidic chip includes a chip cover and a chip substrate, and the chip cover has a cavity channel; the piezoelectric transducer is connected to the chip substrate.

5. A method for introducing microparticles into cells, said method being implemented using a microfluidic device as described in any one of claims 1 to 4, characterized in that, The method includes the following steps: S1, injecting a microparticle solution and a cell suspension into the inlet channel, where the microparticle solution and cell suspension converge and mix before flowing into the main channel; S2, as the mixed liquid flows through the main channel, the gas in at least one microstructure and the mixed liquid in the main channel form a gas-liquid interface at a first gap; S3, controlling the piezoelectric transducer to turn on, the piezoelectric transducer generates sound waves, which enable the gas-liquid interface to vibrate and generate acoustic eddies near the gas-liquid interface; S4, the acoustic eddies cause the cells and microparticles in the mixed liquid to gather and rotate, and the cells undergo reversible deformation under force, generating micropores on the cell membrane, through which the microparticles enter the cell.

6. The method according to claim 5, characterized in that, The method further includes: S5, after a preset loading time, controlling the piezoelectric transducer to close, and the mixed liquid continues to flow in the main channel until it flows out of the outlet channel.

7. The method according to claim 6, characterized in that, The control of the piezoelectric transducer to shut down while the mixed liquid continues to flow in the main channel includes: after a preset shutdown time, restarting the piezoelectric transducer; under the power of the acoustic waves generated by the piezoelectric transducer, acoustic eddies are generated again at multiple gas-liquid interfaces; the acoustic eddies again drive the cells and particles in the mixed liquid to aggregate and rotate; the cells undergo reversible deformation under force, generating micropores on the cell membrane; and the particles enter the cells through the micropores on the cell membrane. After a preset loading time, the control of the piezoelectric transducer to shut down while the mixed liquid continues to flow in the main channel.

8. The method according to claim 7, characterized in that, The method further includes applying an electric field to a microfluidic device, the electric field being used to stimulate and enlarge micropores on the cell membrane.

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