A microfluidic intracellular delivery device and delivery method based on a 3D electroporation system

Abstract

The present invention discloses a microfluidic intracellular delivery device and delivery method based on a 3D electroporation system, relates to the field of intracellular delivery technology, including a 3D electroporation microfluidic chip, a power supply and a pressure pump; the 3D electroporation microfluidic chip is connected to the power supply and the pressure pump respectively; the 3D electroporation microfluidic chip includes a channel import and export and a contraction channel alternately distributed in the middle, the channel import and export are respectively located at the import and export of the chip, for cells to enter the chip and collect cells, and the contraction channel alternately distributed in the middle is a channel with different inner diameters alternately distributed; a power supply is used to provide power to the 3D electroporation microfluidic chip so that an electric field is generated inside the channel, and a pressure pump is used to transport cells through the chip and collect at different flow rates. The present invention can efficiently complete the high-activity delivery of exogenous substances to different types of cells, has excellent delivery efficiency and biocompatibility, and effectively solves the problems of complex operating procedures and large cytotoxicity of current intracellular delivery methods.

Description

Technical Field

[0001] The present invention relates to the field of intracellular delivery technology, and more particularly to a microfluidic intracellular delivery device and delivery method based on a 3D electroporation system. Background Art

[0002] In recent years, microfluidics has become a highly promising approach for intracellular delivery, as its channels and internal structures are sized to match those of cells, leveraging microscale effects to enhance the delivery of exogenous substances. The most commonly used microfluidic intracellular delivery methods include cell squeezing, electroporation, fluid shear-assisted, acoustic-wave-assisted, surface plasmon photothermal-assisted, nanopipette-based microfabrication techniques, and combinations of these techniques. These methods typically manipulate cells through squeezing, shear forces, and electric and acoustic fields enhanced by channel-based structures. These forces exert targeted, appropriately sized forces on the cell membrane, creating pores for the entry of exogenous substances and achieving efficient intracellular delivery, typically achieving delivery efficiencies of approximately 80%. These devices are highly favored for their efficiency, safety, and rapidity. However, flow-based system design must be tailored to cell sizes, which in some cases can lead to contamination from channel clogging and cell fragmentation. Moreover, all of the above methods require devices made with complex micro-nanofabrication technology to achieve the function of intracellular delivery, and the cell processing flux is not high. The pipeline or micro-nano configuration cannot escape the limitations of two-dimensional planar microfabrication, and the design flexibility at the microscale level is not high.

[0003] 3D printing-based microfluidic manufacturing technology has the potential to revolutionize the current status of microfluidic technology in the field of intracellular delivery through distributed organization and practical open source, rapid prototyping methods. For example, a Lego-like The electroporation system, which delivered a green fluorescent protein (GFP) plasmid and a CRISPR-Cas9 plasmid to three-dimensionally cultured HeLa cells, achieved delivery efficiency and cell viability three times higher than that of conventional desktop 3D electroporators, and also verified successful gene editing in HEK293 cells. However, the assembly strategy adds additional steps, and wear and tear during the assembly process can cause unnecessary damage. Therefore, a one-piece 3D-printed microfluidic intracellular delivery device is expected to improve the problems and shortcomings of the assembly process.

[0004] In the field of tumor immunotherapy, T cell gene editing strategies, which modify the expression of chimeric antigen receptors (CARs) to direct T cells to produce cytotoxicity against tumor cells, hold enormous potential. Among gene editing tools, the CRISPR-Cas system is the most disruptive genome editing technology, particularly the Cas9 / ribonucleoprotein complex (RNP). By enabling precise and efficient gene editing while minimizing off-target effects and avoiding unwanted plasmid integration, it is playing an increasingly important role in optimizing CAR-T cells for clinical applications. Although genome editing tools based on the CRISPR / Cas nucleoprotein complex have been widely used for gene editing and gene regulation in cellular settings, inherent limitations of this system, such as the large molecular size of the Cas9 protein complex, limit its further application. Therefore, efficient delivery of gene editing tools into cells is a crucial step in the field of gene editing and gene regulation. The development of novel intracellular delivery methods with high delivery efficiency and minimal cell damage has become a research hotspot. Traditionally, commonly used delivery methods include viral vectors, liposome- or polymer-induced endocytosis, and electrofection.

[0005] Intracellular delivery of CRISPR-Cas plasmids or Cas9 / RNPs integrated into adenoviruses has gained widespread favor due to its exceptionally high efficiency and safety, making it the most commonly used technique for genetic modification. However, concerns remain regarding the potential for immune responses induced by the virus itself and the potential for random integration of viral genes into the host genome, which can disrupt cellular genomic stability. Furthermore, this approach is time-consuming and complex, making it unsuitable for rapid, individualized therapies in the clinical setting. Furthermore, the delivery efficiency of this approach often depends on the cell type and requires rigorous, complex, and standardized experimental environments, which can be detrimental to the normal growth of immune cells (such as primary T cells). Recently, non-viral delivery systems, such as electroporation, have garnered significant attention in clinical research. However, the propensity for cell damage during electroporation and the requirement for specialized equipment have limited their widespread application. Furthermore, to minimize cell damage during electroporation, an additional increase in the number of cells required has hindered advancements in tumor therapy.

[0006] Therefore, how to provide an intracellular delivery method with high efficiency and high cell survival rate that is simple to operate, short in time and high in scalability is an urgent problem that needs to be solved by those skilled in the art. Summary of the Invention

[0007] In view of this, the present invention provides a microfluidic intracellular delivery device and delivery method based on a 3D electroporation system, which can greatly improve the parallel processing capacity of cells by designing multiple radial channels in the channel. Each channel has an alternating distribution of contraction structures, and the narrow geometric structure of the channel is used to achieve local amplification of the electric field strength, thereby realizing the intracellular delivery process. At the same time, the multiple contraction structures in the alternating distribution pattern achieve multiple enhancements of the electric field, thereby replacing the pulse configuration in the traditional electroporator power supply. This method can increase the normal transmembrane potential of the cell (-20 to 200mV) to a transmembrane potential with high permeability (0.2 to 1V) by simply applying a low DC voltage (50 to 200V) and using the narrowed channel structure to achieve a locally enhanced electric field. At the same time, multiple locally enhanced electric fields cause the transmembrane potential of the cell membrane to change multiple times, causing pores in the cell membrane, so that exogenous substances enter the cell to achieve rapid and efficient intracellular delivery. Notably, this method achieves high-throughput (10 million cells per minute) and high-efficiency (approximately 93%) delivery of biomolecules (proteins, plasmids) and functional materials (quantum dots), while maintaining a high cell survival rate (approximately 90%). Furthermore, this technology is widely applicable to a variety of different cell types, including tumor cells, immune cells, adherent cells, suspension cells, and stem cells.

[0008] The present invention efficiently delivered CRISPR-Cas9 plasmids and Cas9 / RNP complexes into A549-mCherry cells and primary T cells, achieving a delivery efficiency of up to 83% while maintaining cell viability at approximately 80%. Characterization and analysis of the percentage of cell fluorescence before and after delivery quickly verified the effectiveness of the system in gene editing operations. The present invention offers high throughput, high efficiency, and minimal damage, offering significant potential for rapid cell-based immunotherapy in clinical settings.

[0009] In order to achieve the above object, the present invention adopts the following technical solutions:

[0010] A microfluidic intracellular delivery device based on a 3D electroporation system includes a 3D electroporation microfluidic chip, a power supply, and a pressure pump; the 3D electroporation microfluidic chip is connected to the power supply and the pressure pump respectively through Tygon hoses, and the three are sequentially integrated and assembled for use;

[0011] The 3D electroporation microfluidic chip includes a methyl methacrylate photopolymer resin and a stainless steel tube. The methyl methacrylate photopolymer resin has a plurality of channels arranged in a three-dimensional pattern. The stainless steel tube corresponds to the position of the channel inlet and outlet. The channels are provided with alternating contraction structures, specifically including the channel inlet and outlet and the intermediate contraction channels. The channel inlet and outlet are located at the inlet and outlet of the chip, respectively, for cell entry and collection. The alternating contraction channels are channels with alternating inner diameters. They are used to generate a high-intensity electric field to create pores in the cell membranes of cells flowing through.

[0012] A power supply, used to provide power to the 3D electroporation microfluidic chip to generate an electric field inside the channel, comprising a DC power supply and a stainless steel tube, wherein the DC power supply is connected to the portion of the stainless steel tube outside the chip;

[0013] A pressure pump is used to transport cells through the chip at different flow rates and collect them.

[0014] Optionally, the 3D electroporation microfluidic chip specifically includes a main inlet, a dispersion hub, an electric field enhancement channel area, an electric field reduction channel area, and a collection hub. The inlet and outlet are symmetrically arranged on both sides of the electric field enhancement channel area, and the power supply is connected to the chip through stainless steel pipes at the inlet and outlet.

[0015] Optionally, the size specifications of the contraction structure channel in the 3D electroporation microfluidic chip are: there are 12 of them, the diameter of a single conventional channel is 0.6 mm, the diameter of the contraction structure channel is 0.3 mm, and the size of the inlet and outlet channels of the 3D electroporation microfluidic chip is: a circular hole with a diameter of 0.8 mm, and the inlet and outlet are correspondingly adhered to the stainless steel pipe.

[0016] Optionally, the dimensions of the stainless steel pipe are: length 10 mm, outer diameter 0.75 mm, inner diameter 0.3 mm, and the stainless steel pipe is connected to a power supply and a pressure pump.

[0017] Optionally, the preparation process of the 3D electroporation microfluidic chip includes:

[0018] First, the 3D electroporation microfluidic chip model, including the channel layout and dimensional parameters, was designed using the Cinema4D R20 software. The model was then uploaded to the Formlabs form 3 3D printer using the Preform software and printing began. After printing, the chip was dismantled and cleaned, and finally UV-cured for 10 minutes to create the 3D electroporation microfluidic chip. Next, a steel pipe of the required size was inserted at each end of the 3D electroporation microfluidic chip to connect the two poles of the power supply and the flexible tube used to transport cells. An electroporation area was designed in the middle of the chip's channel, where the cell suspension flowed in at a constant flow rate from the inlet and was collected at the outlet.

[0019] An intracellular delivery method based on a 3D electroporation system, comprising:

[0020] (1) mixing a cell suspension containing cells with an exogenous substance to ensure that the exogenous substance and the cells are in contact with each other;

[0021] (2) Before delivery, the 3D electroporation microfluidic chip channel was wetted with phosphate buffer solution, and the cells were suspended in phosphate buffer solution containing fluorescein isothiocyanate-labeled dextran;

[0022] (3) The mixture was fed into the inlet of the 3D electroporation microfluidic chip at a constant flow rate and collected at the outlet. To optimize the intracellular delivery parameters, the flow rate was set between 50 and 400 μl / min, and the cell density was maintained at (1-2)×10 5 cells / ml;

[0023] (4) At the same time, a voltage power supply is used to generate a constant DC voltage between the sample and the outlet reservoir, and the voltage value is set in the range of 50-200V;

[0024] (5) A pressure pump is used to supply the cell mixture at a constant flow rate from the inlet of the 3D electroporation microfluidic chip, transport it into the 3D electroporation microfluidic chip system, and finally collect it in a centrifuge tube.

[0025] Optionally, also include:

[0026] (1) centrifuging and washing the dextran molecules with phosphate buffer solution, and then placing them in a cell culture tank under the following culture conditions: culturing at 37°C and 5% CO2 for 12 hours to ensure that the cells adhere to the hollow needle array and spread;

[0027] (2) Characterization and data processing were then performed using an inverted microscope and flow cytometer.

[0028] Through the above technical solution, it can be known that compared with the prior art, the present invention discloses a microfluidic intracellular delivery device and delivery method based on a 3D electroporation system, including a 3D electroporation microfluidic chip, a power supply and a pressure pump; the 3D electroporation microfluidic chip is connected to the power supply and the pressure pump respectively, and the three are integrated and assembled in sequence; the 3D electroporation microfluidic chip includes channels at both ends and contraction channels alternately distributed in the middle, the channels at both ends are respectively located at the inlet and outlet of the chip for cells to enter the chip and collect cells, and the contraction channels alternately distributed in the middle are channels with different inner diameters alternately distributed; a power supply is used to provide power to the 3D electroporation microfluidic chip so that an electric field is generated inside the channel, including a DC power supply and a stainless steel pipe, the DC power supply is connected to the chip outer part of the stainless steel pipe; a pressure pump is used to transport cells through the chip and collect at different flow rates. The present invention can efficiently complete the high-activity delivery of exogenous substances to different types of cells, has excellent delivery efficiency and biocompatibility, and effectively solves the problems of complex operation procedures, low flux and high cytotoxicity of current intracellular delivery methods. BRIEF DESCRIPTION OF THE DRAWINGS

[0029] In order to more clearly illustrate the embodiments of the present invention or the technical solutions in the prior art, the following briefly introduces the drawings required for use in the embodiments or the description of the prior art. Obviously, the drawings described below are merely embodiments of the present invention. For ordinary technicians in this field, other drawings can be obtained based on the provided drawings without paying any creative work.

[0030] Figure 1 Schematic diagram of the intracellular delivery device structure of the 3D electroporation system provided by the present invention;

[0031] Figure 2 is a fluorescent image of the delivered cells in the present invention;

[0032] Figure 3 This is a graph showing the intracellular delivery performance of different inner diameter ratios in the present invention;

[0033] Figure 4 is a graph showing the intracellular delivery performance of the present invention at different contraction numbers;

[0034] Figure 5 is a graph showing the intracellular delivery performance at different voltages and flow rates in the present invention;

[0035] Figure 6 Graph showing the efficiency and cell activity of delivering FITC-dextran molecules of different molecular weights into cells via the intracellular delivery device of the present invention;

[0036] Figure 7 is a graph showing the intracellular delivery performance of the intracellular delivery device of the present invention for different types of cells;

[0037] In the figure, 1-main entrance, 2-dispersion hub, 3-electric field enhancement channel area, 4-electric field reduction channel area, 5-collection hub. DETAILED DESCRIPTION

[0038] The following will clearly and completely describe the technical solutions in the embodiments of the present invention in conjunction with the accompanying drawings. Obviously, the described embodiments are only part of the embodiments of the present invention, not all of the embodiments. Based on the embodiments of the present invention, all other embodiments obtained by ordinary technicians in this field without making creative efforts are within the scope of protection of the present invention.

[0039] The embodiment of the present invention discloses a microfluidic intracellular delivery device based on a 3D electroporation system, comprising a 3D electroporation microfluidic chip, a power supply, and a pressure pump; the 3D electroporation microfluidic chip is connected to the power supply and the pressure pump respectively through Tygon hoses, and the three are sequentially integrated and assembled for use;

[0040] The 3D electroporation microfluidic chip includes a methyl methacrylate photopolymer resin and a stainless steel tube. The methyl methacrylate photopolymer resin has a plurality of channels arranged in a three-dimensional pattern. The stainless steel tube corresponds to the position of the channel inlet and outlet. The channels are provided with alternating contraction structures, specifically including channels at both ends and contraction channels alternately distributed in the middle. The channel inlet and outlet are located at the inlet and outlet of the chip, respectively, for cells to enter the chip and collect cells. The alternating contraction channels are channels with alternating inner diameters. They are used to generate a high-intensity electric field to create pores in the cell membranes of cells flowing through.

[0041] A power supply, used to provide power to the 3D electroporation microfluidic chip to generate an electric field inside the channel, comprising a DC power supply and a stainless steel tube, wherein the DC power supply is connected to the portion of the stainless steel tube outside the chip;

[0042] A pressure pump is used to transport cells through the chip at different flow rates and collect them.

[0043] In a specific embodiment, Figure 1As shown in the figure, the 3D electroporation microfluidic chip specifically includes a main inlet 1, a dispersion hub 2, an electric field enhancement channel area 3, an electric field reduction channel area 4, and a collection hub 5. The inlet and outlet are symmetrically arranged on both sides of the electric field enhancement channel area, and the power supply is connected to the chip through stainless steel pipes at the inlet and outlet. The electric field enhancement area radiates like a star to form 12 microfluidic channels (600μm) evenly distributed in three-dimensional space. Each channel has several contraction and narrowing structures. The 12 radially distributed channels converge again at the same outlet at the bottom of the chip. This design fully utilizes the advantages of flexible three-dimensional space layout while meeting the requirements of high throughput and uniformity of pipeline flow resistance. The single inlet and outlet design also greatly facilitates cell perfusion and collection during electroporation.

[0044] As a mixture of exogenous substances and cells flows through this 3D electroporation microfluidic chip, the narrow structures (300 μm) create a localized electric field enhancement, causing the transmembrane potential (TMP) to exceed the electroporation threshold. This allows contactless electroporation of the cells passing through, enabling high-efficiency, high-throughput, and minimally damaging intracellular delivery. Fine steel needles, each connected to a DC power supply, are inserted at either end of the chip. The main reaction area comprises 12 microfluidic channels, each designed with eight alternating narrow structures with a width-to-narrow ratio of 2 (600:300 μm). At a given flow rate (controlled by a multi-channel microfluidic precision controller / pressure pump), the mixture of exogenous substances and cells achieves intracellular delivery as it flows through the channels. The alternating narrow structures within each channel replace the complex circuitry required to achieve the pulsed action of conventional commercial electroporators. As cells flow through the channel, they experience multiple, dramatic fluctuations in the ambient electric field, similar to the effects of a pulsed electric field, enabling membrane perforation and exogenous substance delivery.

[0045] The preparation of the cell suspension solution, the intelligent fully automatic fluorescence microscopy system and the experimental procedures in the examples are all common knowledge in the art. The model of the fluorescence microscopy system is Ti-E, Nikon, Tokyo, Japan.

[0046] In a specific embodiment, the size specifications of the contraction structure channel in the 3D electroporation microfluidic chip are: there are 12, the diameter of a single conventional channel is 0.6 mm, the diameter of the contraction structure channel is 0.3 mm, and the inlet and outlet channel dimensions of the 3D electroporation microfluidic chip are: circular holes with a diameter of 0.8 mm, and the inlet and outlet are correspondingly adhered to the stainless steel pipe.

[0047] In a specific embodiment, the dimensions of the stainless steel pipe are: length 10 mm, outer diameter 0.75 mm, inner diameter 0.3 mm, and the stainless steel pipe is connected to a power supply and a pressure pump.

[0048] In a specific embodiment, this embodiment provides a 3D electroporation microfluidic chip channel design with a diameter larger than the cell diameter to ensure that the cells can pass through the narrow structure intact, avoiding cell damage problems caused by cell accumulation and channel blockage, thereby effectively improving the flux and activity of intracellular delivery. It is worth noting that compared with existing membrane destruction technologies such as cell extrusion, the chip system has significant advantages. The former may cause serious cell damage and blockage when cells interact with the tube wall, while the wide microchannels of the chip ensure that the cells flow freely through the channel, thereby maintaining the stability of the system and greatly improving the work efficiency and the flux of a single transfection. The results show that when the cells flow through the channel at high speed, they hardly collide directly with the tube wall.

[0049] The preparation process of 3D electroporation microfluidic chip includes:

[0050] Preparation of 3D electroporation microfluidic chip

[0051] The chip portion of the intracellular delivery device was fabricated using 3D printing technology. First, a chip model was designed using Cinema4D R20, including parameters such as channel layout and dimensions. The model was then uploaded to a Formlabs form 3 3D printer using the Preform software and printing began. After printing, the chip was disassembled, cleaned, and finally UV-cured for 10 minutes. Next, a suitably sized steel pipe was inserted at each end of the chip to connect the two terminals of the power supply and the flexible tube used to transport the cells. The central portion of the chip's channels houses an electroporation zone, where the cell suspension flows at a constant flow rate from the inlet and is collected at the outlet. The narrowing of the channel's inner diameter (reduced from 600 to 300 microns) is designed to achieve a high electric field at low voltage, creating a highly efficient electroporation zone. The 12 channels with this function are arranged in three dimensions to increase intracellular delivery throughput, prevent channel clogging, and enhance cell activity.

[0052] (2) Assembly and optimization of intracellular delivery devices

[0053] The intracellular delivery device primarily consists of a microfluidic chip, a DC power supply, and a pressure pump. The DC power supply provides a stable, controllable voltage. By adjusting the DC power supply voltage, the output voltage ranges from 0 to 210V, enabling gradient control of the electric field intensity within the chip channel. The flow rate of the cell suspension is controlled by the pressure pump, which outputs a pressure range of 50 to 2000 mbar and a flow rate of 0 to 4 mL / min.

[0054] The microfluidic intracellular delivery device based on the 3D electroporation system prepared by 3D printing was used in the intracellular delivery experiment, specifically:

[0055] (1) mixing a cell suspension containing cells with an exogenous substance to ensure that the exogenous substance and the cells are in contact with each other;

[0056] (2) Before the delivery experiment, the chip channel was moistened with phosphate buffered saline (PBS). The cells were suspended in a PBS solution containing fluorescein isothiocyanate-labeled dextran (FITC-dextran). The concentrations of FITC-dextran were 100 μg / ml, 300 μg / ml, 500 μg / ml, and 1000 μg / ml, respectively.

[0057] (3) The mixture was fed from the chip inlet at a constant flow rate and collected at the outlet. To optimize the intracellular delivery parameters, the flow rate was set between 50 and 400 μl / min and the cell density was maintained at (1-2)×10 5 cells / ml.

[0058] (4) At the same time, a voltage source (2400 Source Meter, Keithley) was used to generate a constant direct current (DC) voltage between the sample and the outlet reservoir, and the voltage value was set to 110 V.

[0059] (5) A pressure pump is used to supply the cell mixture at a constant flow rate from the chip inlet to the chip system. When the cells flow through the narrow part of the channel, the field intensity is higher than the electroporation threshold, and electroporation occurs. The alternating distribution structure produces a pulse effect that can only be produced by a traditional power supply complex circuit, further increasing the probability of the electric field generating pores on the cell membrane, thereby increasing the chances of exogenous substances entering the cell and ensuring a higher delivery efficiency. Finally, the cells are collected in a centrifuge tube, centrifuged, and the FITC-dextran molecules are washed with PBS. Finally, the collected cells are placed in a 37°C incubator for 12 hours, and then characterized and data processed using instruments such as an inverted microscope and a flow cytometer. The results are as follows: Figure 2 As shown, cells delivered by the intracellular delivery device have obvious fluorescence, indicating successful intracellular delivery.

[0060] The intracellular delivery device described in the examples was used to study the delivery of biomolecules to cells:

[0061] (1) Effect of the channel inner diameter ratio of the intracellular delivery device on the delivery efficiency of FITC-dextran molecules

[0062] This embodiment designs the effect of the inner diameter ratio of the delivery structure in different pipes (that is, the diameter ratio of the conventional pipe and the narrow channel) on the delivery efficiency. Based on the intracellular delivery device, a series of delivery experiments and result characterizations were carried out on human non-small cell lung cancer cells (A549) using the FITC-dextran molecule with a molecular weight of 70kDa as a model molecule. The results showed that: compared with the pipe without a narrow structure (control group), there was a significant fluorescence enhancement in the cells after the cells were electroporated through the chip with narrow contraction, indicating that the FITC-dextran molecules were able to enter the interior of the cells through the intracellular delivery device. The relationship between the channel width-to-inner diameter ratio and the delivery performance of the intracellular delivery device was analyzed: the four inner diameter ratios were 1, 1.5, 2 and 2.5, respectively. At the same time, the number of channel narrowing structures (#5), the applied voltage (100V), and the flow rate (2mL / min) remained unchanged. The results are as follows. Figure 3 As shown, under the conditions of the number of channel narrowing structures, applied voltage and flow, the delivery efficiency tends to increase with the increase of the ratio. When the channel width to inner diameter ratio is 2.5, the delivery efficiency reaches 80%. Because as the channel width to inner diameter ratio increases, the amplitude of the electric field change becomes larger, resulting in a stronger stimulation to the cell membrane, which may cause more pores to form on the cell membrane, and the probability that exogenous substances can enter the cell is greater. However, when the amplitude of the change in the electric field intensity acting on the cell is too large, it will cause irreversible damage to the cell membrane, that is, reduce the activity of the cell. It is necessary to choose a method that has a higher delivery efficiency while maintaining higher cell activity. For the cell activity experiment, CCK-8 reagent was used to evaluate the activity of the cells after passing through the intracellular delivery device, and the statistical results are as follows. Figure 3 As shown in Figure 2, as the channel width-to-inner diameter ratio increases, cell activity decreases. The results show that an inner diameter ratio of 2 is the most suitable, as it maintains high efficiency while maintaining cell activity exceeding 80%.

[0063] (2) Effect of the amount of channel narrowing in the intracellular delivery device on the delivery efficiency of FITC-dextran molecules

[0064] This example further evaluated and analyzed the effect of the number of channel narrowing structures on the intracellular delivery performance of the intracellular delivery device. While maintaining the applied voltage (100V), flow rate (2mL / min) and channel width-to-inner diameter ratio of 2, the number of single channel narrowing structures was increased from 0 to 10. The corresponding delivery efficiency and cell viability results are shown in Figure 2. Figure 4The experimental statistical results are consistent with the hypothesis. As the number of constriction structures increases, the efficiency of intracellular delivery gradually increases. The chip with 10 constriction structures in a single channel exhibits the highest delivery efficiency (approximately 83%). However, because the number of channel constriction structures directly determines the frequency with which cells experience electric field fluctuations, excessively high frequencies may cause excessive cell damage. Therefore, it is considered more appropriate to have 8 constriction structures per channel.

[0065] (3) Effect of the electric field of the intracellular delivery device channel on the delivery efficiency of FITC-dextran molecules

[0066] The external applied voltage is also an important parameter that can affect the delivery efficiency. In this embodiment, the designed micron-scale channel can generate a sufficiently high electric field (about 81.6 kV cm) at low voltage. -1 Using a lower DC voltage can not only reduce the cost of the power supply and avoid unnecessary electrolytic side reactions, but also achieve miniaturization of the device. In order to determine the voltage value that can achieve the best intracellular delivery performance of the device, the voltage values ​​applied at the inlet and outlet were set to 0V, 50V, 80V, 110V, 140V, 170V, and 200V by adjusting the voltage. The results are shown in Figure 2. Figure 5 As shown in the figure, when the applied voltage value increased to 80 V and 200 V, the transport efficiency reached over 80% and 95%, respectively. Higher voltage conditions caused some cell damage, so the most suitable applied voltage was selected as 110 V in subsequent experiments.

[0067] The effect of the coupling relationship between the external electric field and the constriction structure on the efficiency of intracellular delivery was further analyzed. To this end, the following group experiments were set up, including endocytosis group, chip without structure and no voltage group, chip with structure and no voltage group, chip without structure and voltage group, and chip with structure and voltage group. Through flow cytometric statistical analysis of transfection efficiency, it was found that ( Figure 5 ). Compared to the endocytosis group, weak transfection signals were observed in both the no-structure, no-voltage and the with-structure, no-voltage groups, indicating either insufficient FITC-dextran endocytosis or a surface binding event. Furthermore, when both structure and voltage were present, intracellular delivery performance was significantly improved, exceeding 90% efficiency, demonstrating that this combined structure-voltage intracellular delivery model maximizes electroporation benefits under conditions of local electric field fluctuations.

[0068] (4) Effect of flow rate in the intracellular delivery device channel on the delivery efficiency of FITC-dextran molecules

[0069] Obviously, the delivery flux depends to a great extent on the appropriate flow rate. The analysis of the relationship between flow rate and delivery efficiency shows that the delivery efficiency shows a trend of first increasing and then decreasing with the increase of flow rate, reaching 85% under the condition of 2 mL / min. Figure 5 These results indicate that the shear force generated by the optimal flow rate facilitates cell deformation as it passes through the channel, ultimately promoting the formation of cell membrane pores and improving delivery efficiency. Conversely, excessively fast flow rates limit the residence time of cells and external substances in the channel, resulting in reduced delivery efficiency. Therefore, ideal intracellular delivery flux can be achieved at an appropriate flow rate.

[0070] (5) Effect of intracellular delivery devices on the delivery efficiency of FITC-dextran molecules of different molecular weights:

[0071] In order to further verify the universality of the intracellular delivery device, we first demonstrated that molecules of different molecular weights can be delivered intracellularly through the system. First, we tested the delivery of FITC-dextran molecules with molecular weights between 4-500 kDa into A549 cells under the same conditions. Figure 6 The statistical results of the delivery efficiency shown show that the delivery efficiency can reach up to 95%, and the intracellular delivery device can still show a high efficiency of 89% when facing the delivery of large molecules. This shows that the system can efficiently deliver biological molecules of different molecular weights into cells, that is, it is not limited by the size of exogenous molecules. Not only that, compared with other delivery methods, the intracellular delivery device is suitable for different types of cells and has a wide range of applicability. In order to evaluate the delivery ability of the intracellular delivery device in different cell types, FITC-dextran molecules with a molecular weight of 70kDa were delivered into HEK293T, A549, HepG2, Jurkat, U87 and primary T cells through the system, and the delivery efficiency and cell activity were statistically analyzed. The results showed that the FITC-dextran molecule with a molecular weight of 70kDa was successfully delivered to these cells, and good delivery efficiency and cell activity were obtained, proving that the intracellular delivery device has significant versatility ( Figure 7 ).

[0072] Extensive comparisons indicate that the intracellular delivery device, with its wider channel structure, prevents cell damage from squeezing or other mechanical forces. This demonstrates that the device we designed and fabricated can serve as a suitable alternative to current intracellular delivery methods, which are hampered by complex operational procedures, different formulation requirements for different cell types, and cytotoxicity. More importantly, these results demonstrate that the intracellular delivery device exhibits excellent intracellular delivery performance for diverse cell types, particularly immune cells, giving it significant potential for development in cellular immunotherapy and other important applications.

[0073] These results demonstrate that the intracellular delivery device exhibits excellent intracellular delivery performance across diverse cell types, giving it significant potential for development in cellular immunotherapy and other important applications. They also demonstrate that the device we designed and fabricated can serve as a suitable alternative to current intracellular delivery methods, which are hampered by complex operational procedures, the need for different formulations for different cell types, and cytotoxicity.

[0074] The above description of the disclosed embodiments is intended to enable one skilled in the art to implement or use the present invention. Various modifications to these embodiments will be readily apparent to one skilled in the art, and the general principles defined herein may be implemented in other embodiments without departing from the spirit or scope of the present invention. Therefore, the present invention is not limited to the embodiments shown herein but is intended to conform to the widest scope consistent with the principles and novel features disclosed herein.

Claims

1. A microfluidic intracellular electroporation system based on 3D electroporation The delivery device is characterized by comprising a 3D electroporation microfluidic chip, a power supply, and a pressure pump; the 3D electroporation microfluidic chip is connected to the power supply and the pressure pump respectively through Tygon hoses, and the three are sequentially integrated and assembled for use; The 3D electroporation microfluidic chip includes a methyl methacrylate photopolymer resin and a stainless steel tube. The methyl methacrylate photopolymer resin has a plurality of channels arranged in a three-dimensional pattern. The stainless steel tube corresponds to the position of the channel inlet and outlet. The channels are provided with alternating contraction structures, specifically including the channel inlet and outlet and the intermediate contraction channels. The channel inlet and outlet are located at the inlet and outlet of the chip, respectively, for cell entry and collection. The alternating contraction channels are channels with alternating inner diameters. They are used to generate a high-intensity electric field to create pores in the cell membranes of cells flowing through. A power supply, used to provide power to the 3D electroporation microfluidic chip to generate an electric field inside the channel, comprising a DC power supply and a stainless steel tube, wherein the DC power supply is connected to the portion of the stainless steel tube outside the chip; A pressure pump is used to transport cells through the chip at different flow rates and collect them.

2. A microfluidic intracellular delivery device based on a 3D electroporation system according to claim 1, characterized in that: The 3D electroporation microfluidic chip specifically includes a main inlet, a dispersion hub, an electric field enhancement channel area, an electric field reduction channel area, and a collection hub. The inlet and outlet are symmetrically arranged on both sides of the electric field enhancement channel area, and the power supply is connected to the chip through stainless steel pipes at the inlet and outlet.

3. The microfluidic intracellular delivery device based on a 3D electroporation system according to claim 2, characterized in that: The dimensions of the constriction structure channels in the 3D electroporation microfluidic chip are as follows: there are 12 of them, with a single conventional channel having a diameter of 0.6 mm and a constriction structure channel having a diameter of 0.3 mm. The inlet and outlet channels of the 3D electroporation microfluidic chip are circular holes with a diameter of 0.8 mm, and the inlet and outlet are adhered to the stainless steel pipes.

4. The microfluidic intracellular delivery device based on a 3D electroporation system according to claim 1, characterized in that: The dimensions of the stainless steel pipe are: 10 mm in length, 0.75 mm in outer diameter, and 0.3 mm in inner diameter. The stainless steel pipe is connected to a power supply and a pressure pump.

5. The microfluidic intracellular delivery device based on a 3D electroporation system according to claim 1, characterized in that: The preparation process of the 3D electroporation microfluidic chip includes: First, the 3D electroporation microfluidic chip model, including the channel layout and dimensional parameters, was designed using the Cinema4D R20 software. The model was then uploaded to the Formlabs form 3 3D printer using the Preform software and printing began. After printing, the chip was dismantled and cleaned, and finally UV-cured for 10 minutes to create the 3D electroporation microfluidic chip. Next, a steel pipe of the required size was inserted at each end of the 3D electroporation microfluidic chip to connect the two poles of the power supply and the flexible tube used to transport cells. An electroporation area was designed in the middle of the chip's channel, where the cell suspension flowed in at a constant flow rate from the inlet and was collected at the outlet.

6. An intracellular delivery method based on a 3D electroporation system, characterized in that: A microfluidic intracellular delivery device based on a 3D electroporation system according to any one of claims 1 to 5, comprising: (1) mixing a cell suspension containing cells with an exogenous substance to ensure that the exogenous substance and the cells are in contact with each other; (2) Before delivery, the 3D electroporation microfluidic chip channel was wetted with phosphate buffer solution, and the cells were suspended in phosphate buffer solution containing fluorescein isothiocyanate-labeled dextran; (3) The mixture was fed into the inlet of the 3D electroporation microfluidic chip at a constant flow rate and collected at the outlet. To optimize the intracellular delivery parameters, the flow rate was set between 50 and 400 μl / min, and the cell density was maintained at (1-2)×10 5 cells / ml; (4) At the same time, a voltage power supply is used to generate a constant DC voltage between the sample and the outlet reservoir, with the voltage value setting range being 50-200 V; (5) A pressure pump is used to supply the cell mixture at a constant flow rate from the inlet of the 3D electroporation microfluidic chip, transport it into the 3D electroporation microfluidic chip system, and finally collect it in a centrifuge tube.

7. The intracellular delivery method based on a 3D electroporation system according to claim 6, characterized in that: Also includes: (1) centrifuging and washing the dextran molecules with phosphate buffer solution, and then placing them in a cell culture tank under the following culture conditions: culturing at 37°C and 5% CO2 for 12 hours to ensure that the cells adhere to the hollow needle array and spread; (2) Characterization and data processing were then performed using an inverted microscope and flow cytometer.

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