Microfluidic chip and microfluidic chip-based biological particle separation device

Abstract

The application provides a microfluidic chip and a biological particle separation device based on the microfluidic chip, which comprises a main flow channel, a plurality of bypass flow channels and an electrode flow channel. The electrode flow channel is arranged on both sides of the side wall of the main flow channel and forms a plurality of electrode zones along the extension direction of the main flow channel, so as to provide a non-uniform electric field for the main flow channel; the plurality of bypass flow channels are sequentially communicated with the main flow channel along the extension direction of the main flow channel, so that the electrode zones and the bypass flow channels are alternately and spacedly arranged along the extension direction of the main flow channel. Thus, after the sample liquid is injected into the main flow channel, different target biological particles in the sample liquid can produce different lateral deviations under the dielectrophoresis force of one or more electrode zones, so that the target biological particles can flow out from the liquid outlet of the corresponding bypass flow channel or the main flow channel, and the efficient separation of biological particles of multiple sizes and types is realized.

Description

TECHNICAL FIELD

[0001] The present application relates to the field of microfluidic technology, and in particular to a microfluidic chip and a biological particle separation device based on the microfluidic chip. BACKGROUND

[0002] In the field of biomedicine, etc., accurate separation of multiple particles is crucial for the processing and analysis of complex biological samples. Whether it is clinical diagnosis or biomedical research, blood, body fluid and other samples contain a large number of heterogeneous particles, and the efficient separation of these particles is the basis for accurate analysis and diagnosis.

[0003] Dielectrophoresis microfluidic separation technology utilizes the dielectric properties of particles under the action of a non-uniform electric field to realize the movement and separation of particles in the electric field, making it one of the main ways to realize particle separation. Compared with traditional labeling separation methods, dielectrophoresis microfluidic separation technology does not require any labeling of particles, avoiding the interference of the labeling process on the natural state of the particles. However, the application of existing dielectrophoresis microfluidic separation technology is mostly based on the separation of single particle characteristics (such as size or dielectric properties), and usually only binary separation can be achieved, i.e. the sample is divided into two categories. For complex biological samples containing multiple sizes and types of particles, the separation efficiency and accuracy of this method are limited.

[0004] Therefore, how to provide a microfluidic chip and a biological particle separation device based on the microfluidic chip, which can overcome the limitations of existing microfluidic dielectrophoresis technology in processing complex biological samples, and realize efficient separation of multiple sizes and types of biological particles, has become one of the technical problems that personnel in the field need to solve. SUMMARY

[0005] The present application aims to at least partially solve one of the technical problems in the related art.

[0006] To this end, the first purpose of the present application is to provide a microfluidic chip and a biological particle separation device based on the microfluidic chip, which can overcome the limitations of existing microfluidic dielectrophoresis technology in processing complex biological samples, and realize efficient separation of multiple sizes and types of biological particles.

[0007] To achieve the above-mentioned purpose, the first aspect of the present application provides a microfluidic chip, comprising a main flow channel, a plurality of bypass flow channels and an electrode flow channel; wherein,

[0008] One end of the main flow channel is connected with a sample inlet and a plurality of sheath liquid inlets, and the opposite end is connected with a plurality of liquid outlets; sample liquid and sheath liquid are injected into the main flow channel after being converged through the sample inlet and the sheath liquid inlet respectively;

[0009] The electrode flow channels are arranged on both sides of the side walls of the main flow channel in the second direction, and form a plurality of electrode zones in the first direction, each of the electrode zones providing a non-uniform electric field for the main flow channel at a corresponding position in the second direction, the first direction being the extension direction of the main flow channel, and the second direction being perpendicular to the extension direction of the main flow channel.

[0010] The plurality of bypass flow channels are sequentially communicated with the main flow channel in the first direction, and each of the bypass flow channels corresponds to each of the electrode zones, and the electrode zones and the bypass flow channels are alternately spaced in the first direction, so that different target biological particles in the sample liquid injected into the main flow channel can be laterally offset in the second direction under the action of dielectrophoresis force of one or more electrode zones, and flow out from the corresponding bypass flow channel or the liquid outlet.

[0011] Optionally, one end of the main flow channel is connected with at least one sample inlet and at least two sheath liquid inlets, and the other end of the main flow channel is connected with at least one liquid outlet.

[0012] Optionally, each of the bypass flow channels comprises a direct current output part, a serpentine flow channel part and a direct current connection part, the direct current connection part communicates the serpentine flow channel part with the main flow channel, and the direct current output part communicates the serpentine flow channel part with a corresponding outlet.

[0013] Optionally, the electrode flow channel comprises a first electrode flow channel and a second electrode flow channel, and the first electrode flow channel and the second electrode flow channel are respectively filled with liquid metal to form the first electrode and the second electrode; different target biological particles in the main flow channel move to the side close to the first electrode in the second direction under the action of dielectrophoresis force of one or more electrode zones.

[0014] Optionally, the first electrode comprises a plurality of first sub-electrodes, and the plurality of first sub-electrodes correspond to the second electrode to form a plurality of electrode zones spaced from each other; wherein the first electrode is a positive electrode, and the second electrode is a negative electrode.

[0015] Optionally, the main flow channel corresponding to each of the electrode zones can be divided into a bypass flow field close to the first electrode and a main flow field away from the first electrode in the second direction, and when the target biological particles enter the bypass flow field from the main flow field under the action of dielectrophoresis force of the electrode zone, the target biological particles will flow out from the bypass flow channel corresponding to the electrode zone.

[0016] Optionally, when the target biological particles in the corresponding electrode zone satisfy the following relationship, the target biological particles will flow out from the corresponding bypass flow channel, i.e.

[0017] d0-d DEP <w b ,

[0018] wherein, d0 is an initial position of the target biological particle along the second direction when entering the electrode region, the d DEP is an offset of the target biological particle along the second direction under the dielectrophoresis force of the electrode region; w b is a width of the bypass flow field along the second direction.

[0019] Optionally, the radius of the target biological particle is less than the width of the bypass flow field along the second direction.

[0020] Optionally, the length and width of the main flow channel corresponding to different electrode regions are not the same, and the length and width of each bypass flow channel correspond to the length and width of the main flow channel corresponding to the electrode region.

[0021] To achieve the above object, the second aspect of the present application provides a biological particle separation device based on a microfluidic chip, comprising the microfluidic chip described in any one of the above.

[0022] The microfluidic chip and the biological particle separation device based on the microfluidic chip provided by the present application at least have the following beneficial effects:

[0023] The present application provides a microfluidic chip and a biological particle separation device based on the microfluidic chip, which comprises a main flow channel, a plurality of bypass flow channels and an electrode flow channel. Wherein, the electrode flow channel is arranged on both sides of the side wall of the main flow channel and forms a plurality of electrode regions along the extension direction of the main flow channel, providing a non-uniform electric field for the main flow channel; the plurality of bypass flow channels are sequentially communicated with the main flow channel in the extension direction of the main flow channel, so that the electrode region and the bypass flow channel can be arranged alternately and spaced along the extension direction of the main flow channel. Different target biological particles in the sample liquid injected into the main flow channel can produce different lateral offsets in the normal direction of the flow direction of the sample liquid under the dielectrophoresis force of one or more electrode regions, so that the target biological particles can flow out of the liquid outlet corresponding to the bypass flow channel or the main flow channel. The present application combines the passive flow field shunting mechanism with the active cumulative dielectrophoresis deflection effect, and realizes the efficient separation of biological particles of various sizes and types.

[0024] Additional aspects and advantages of the present application will be partially given in the following description, partially will become obvious from the following description, or will be understood by the practice of the present application. BRIEF DESCRIPTION OF DRAWINGS

[0025] The above and / or additional aspects and advantages of the present application will become apparent and more readily appreciated from the following description of the embodiments, taken in conjunction with the accompanying drawings, in which:

[0026] Figure 1A structure schematic diagram of a microfluidic chip according to an embodiment of the present application.

[0027] Figure 2 A cross-sectional structure schematic diagram of a microfluidic chip according to an embodiment of the present application.

[0028] Figure 3 A cross-sectional structure schematic diagram of another microfluidic chip according to an embodiment of the present application.

[0029] Figure 4 A structure and flow resistance schematic diagram of a bypass flow channel according to an embodiment of the present application.

[0030] Figure 5 A corresponding relationship curve between a flow resistance coefficient and a flow channel number according to an embodiment of the present application.

[0031] Figure 6 A structure schematic diagram of an inlet channel according to an embodiment of the present application.

[0032] Figure 7 A schematic diagram of a uniform sphere model and a single-shell sphere model according to an embodiment of the present application.

[0033] Figure 8 A deflection simulation schematic diagram of different target particles under the action of different dielectrophoretic forces according to an embodiment of the present application.

[0034] Figure 9 A biological particle separation device verification test result schematic diagram according to an embodiment of the present application.

[0035] 100 main flow channel; 101 sample inlet; 102 sheath liquid inlet; 103 outlet; 200 electrode flow channel; 201 electrode area; 210 first electrode flow channel; 211 first electrode; 220 second electrode flow channel; 221 second electrode; 300 bypass flow channel. DETAILED DESCRIPTION

[0036] The embodiments of the present application are described in detail below, examples of which are shown in the accompanying drawings, wherein the same or similar notations represent the same or similar elements or elements having the same or similar functions throughout. The embodiments described below by referring to the drawings are exemplary and are intended to explain the present application, and cannot be understood as a limitation of the present application.

[0037] The existing multi-particle separation technology generally includes traditional particle separation technology and microfluidic particle separation technology, and the microfluidic separation technology includes label-based microfluidic technology and label-free microfluidic technology.

[0038] Traditional particle separation techniques rely on differences in physical properties such as size, density, etc., and are separated by methods such as centrifugation, filtration, etc. Although these methods can separate different types of particles to some extent, their separation efficiency and accuracy are limited when facing high heterogeneity and multiple types of particles. Especially when the sample contains particles with a wide size range and similar physical properties, such as red blood cells, white blood cells, platelets, and other particles in a complex sample, traditional particle separation techniques often fail to achieve effective separation.

[0039] Label-based microfluidic techniques rely on labeling methods such as fluorescent labeling or magnetic labeling, and separate labeled particles by external fields such as electric fields or magnetic fields. Although labeling techniques can improve the selectivity of separation, they also bring new problems, including complex labeling process, high cost, and possible interference with the natural state and function of biological particles. For particles that are not easy to label or not suitable for labeling, the applicability of traditional microfluidic techniques is also limited.

[0040] Label-free microfluidic techniques use the inherent physical and chemical properties of biological particles to achieve separation, such as operating on the differences in size, density, or dielectric properties of particles. Such methods avoid the interference and complexity of the labeling process, thus maintaining the original biological activity of the sample, simplifying the experimental process, and reducing cost and time consumption. However, the strategy of integrating multiple separation techniques in a single device can improve separation efficiency and selectivity, but the design and manufacturing process is complex, increasing the difficulty of operation and equipment cost. The matching and regulation of these integrated devices under high flow conditions are also complex, limiting their practical application. The strategy of using a single physical mechanism for separation simplifies the design and operation, but it is still difficult to efficiently separate multiple particle types in the same device.

[0041] Based on the above problems, the embodiments of the present application provide a microfluidic chip and a biological particle separation device based on the microfluidic chip, which combines passive flow field splitting mechanism and active accumulation dielectrophoresis deflection effect to achieve efficient separation of multiple particles. At the same time, by optimizing the electrode configuration and flow channel design, the present application can efficiently separate complex biological samples containing different sizes and dielectric properties under simple electric field configuration, providing an efficient, simple, and reliable separation technique for target particles in complex biological samples, with high universality and adaptability.

[0042] According to a first aspect of the present application, a microfluidic chip is provided, such as Figures 1-3As shown, the chip includes a substrate and a flow channel layer, the flow channel layer is formed on the substrate, and a plurality of grooves are etched on the surface close to the substrate side, and the plurality of grooves combine with the substrate to form a main flow channel 100, an electrode flow channel 200 and a plurality of bypass flow channels 300.

[0043] The material of the substrate can be inorganic rigid material, including but not limited to one of monocrystalline silicon, amorphous silicon, glass, quartz and the like; the material of the flow channel layer is organic polymer material, including but not limited to one of epoxy resin, polyurea, polyurethane, polystyrene and polymethyl methacrylate, polydimethylsiloxane (PDMS) and the like. In the embodiment, the material of the substrate is glass, and the material of the flow channel layer is PDMS.

[0044] As shown in Figure 2 and Figure 3 One end (the first end) of the main flow channel 100 is connected with the inlet channel, and the opposite end (the last end) is connected with the outlet channel, so that the sample liquid can flow in the main flow channel 100. Among them, the inlet channel includes a sample inlet 101 and a plurality of sheath inlets 102, the sample liquid is injected into the inlet channel through the sample inlet 101, and the focusing sheath liquid is injected into the inlet channel through the sheath inlet 102, so that the sample liquid and the focusing sheath liquid can be injected into the main flow channel 100 after converging in the inlet channel.

[0045] By connecting one end of the main flow channel 100 with the sample inlet 101 and the plurality of sheath inlets 102, the target biological particles in the sample liquid injected into the main flow channel 100 can flow along the main flow channel 100 under the joint action of the focusing sheath liquid flow field and the sample liquid flow field, and by appropriately adjusting the input flow of the focusing sheath liquid and the sample liquid, the position of the target biological particles entering the main flow channel 100 in the direction perpendicular to the extension direction of the main flow channel 100 can be adaptively adjusted.

[0046] For convenience of description, the extension direction of the main flow channel 100 is set as the first direction, i.e. the X direction in the figure, and the direction perpendicular to the extension direction of the main flow channel 100 is set as the second direction, i.e. the Y direction in the figure, and the orthogonal plane of the X direction and the Y direction is perpendicular to the substrate surface.

[0047] It should be noted that the sample liquid can be a biological particle suspension with a plurality of target biological particles, and the expressions "target biological particle", "target particle" or "biological particle" will be used in this application, and sometimes will be directly omitted as "particle". The above expressions have the same meaning and can be used interchangeably, and the specific meaning of each of them is to refer to the target particle to be separated in the sample liquid.

[0048] The electrode flow channel 200 is arranged on both sides of the side wall of the main flow channel 100 along the second direction, and forms a plurality of discontinuous electrode regions 201 in the first direction, each electrode region 201 provides a non-uniform electric field for the corresponding position of the main flow channel 100 in the second direction, so that a plurality of target biological particles in the main flow channel 100 can be subjected to dielectrophoresis in different electrode regions 201 along the second direction, and then a plurality of target biological particles in the main flow channel 100 are offset along the second direction under the cumulative action of different dielectrophoretic forces.

[0049] Dielectrophoresis is a technology that uses electric field force in a non-uniform electric field to act on polarized particles. In a non-uniform electric field, particles move directionally due to dielectrophoretic force, thereby achieving separation. The dielectrophoretic force experienced by the particles is related to the size of the particles, so different sizes of particles require different dielectrophoretic forces to achieve effective separation. Under the condition that the dielectric properties of the medium and biological particles are determined, the dielectrophoretic force is mainly determined by the non-uniformity of the electric field and the position of the particles in the electric field.

[0050] Different target biological particles in the main flow channel 100 will produce different degrees of offset along the second direction under the action of the dielectrophoretic force of one or more electrode regions 201 according to their size and dielectric properties, and the position of the biological particles in the first direction changes with the flow of the sample liquid and the focusing sheath liquid in the main flow channel 100, thereby causing a plurality of target biological particles to be layered in the orthogonal plane formed by the first direction and the second direction.

[0051] The electrode flow channel 200 can specifically include a first electrode flow channel 210 and a second electrode flow channel 220, and the first electrode flow channel 210 and the second electrode flow channel 220 are respectively filled with liquid metal to form two three-dimensional electrodes, namely the first electrode 211 and the second electrode 221, after being powered on. Among them, the polarity of the first electrode 211 and the second electrode 221 is opposite, such as the first electrode 211 being a positive electrode and the second electrode 221 being a negative electrode, or the first electrode 211 being a negative electrode and the second electrode 221 being a positive electrode. The specific polarity of the first electrode 211 and the second electrode 221 is not limited in the present application. In order to facilitate the description of the subsequent embodiments, the first electrode 211 can be set as a positive electrode and the second electrode 221 can be set as a negative electrode.

[0052] Therefore, the first electrode 211 and the second electrode 221 are arranged on both sides of the side wall of the main flow channel 100 along the second direction, and provide a non-uniform electric field for the main flow channel 100, so that different target biological particles in the main flow channel 100 can be deflected and moved along the second direction to the side close to the first electrode 211 under the action of the dielectrophoretic force of one or more electrode regions 201.

[0053] As an example, the first electrode flow channel 210 can include one main electrode flow channel and multiple sub-electrode flow channels in communication with the main electrode flow channel, and the liquid metal is sequentially injected into the multiple sub-electrode flow channels and sufficiently fills each sub-electrode flow channel to form multiple discontinuous first sub-electrodes on the side of the main flow channel 100 away from the second electrode 221. That is, the first electrode 211 can be composed of multiple first sub-electrodes corresponding to the second electrode 221 at corresponding positions, forming multiple electrode regions 201 spaced apart from each other in the first direction.

[0054] As an example, each sub-electrode flow channel further includes multiple first through-holes, such as 10 μm holes, arranged in the first direction on the side close to the main flow channel 100, each first through-hole communicating each sub-electrode flow channel with the main flow channel 100 to form a first electrode array. Correspondingly, the second electrode flow channel 220 in the corresponding range of each electrode region 201 further includes multiple second through-holes, such as 10 μm holes, arranged in the first direction on the side close to the main flow channel 100, each second through-hole communicating the second electrode flow channel 220 with the main flow channel 100 to form a second electrode array. Among them, the number of first through-holes is less than the number of second through-holes, and the asymmetric hole structure composed of first through-holes and second through-holes, whereby the first electrode array and the second electrode array of each electrode region 201 can form a non-uniform electric field required for dielectrophoresis at the main flow channel 100 when the first electrode 211 and the second electrode 221 are powered.

[0055] It should be noted that the liquid metal includes but is not limited to metals such as indium, tin, cadmium, bismuth, lead and their alloy liquids such as liquid gallium and gallium-indium alloy, and is formed by cooling and solidification after being injected into the electrode flow channel 200 by hand pushing or other mechanical means. The liquid metal is injected into the electrode flow channel 200 in liquid state at a relatively high operating temperature, and then the microfluidic chip environment temperature is reduced to room temperature or use temperature, so that the liquid metal solidifies.

[0056] In addition, other non-electric physical field methods can also be used to achieve active displacement of particles in the main flow channel 100, such as surface acoustic wave field, magnetic field, light field, etc., as long as the spatial distribution of the active physical field is set to have appropriate gradient changes.

[0057] The plurality of bypass flow channels 300 are sequentially communicated with the main flow channel 100 along the first direction, and each bypass flow channel 300 corresponds to each electrode area 201 one by one, and is alternately spaced along the first direction, so that different target biological particles in the sample liquid injected into the main flow channel 100 can produce different lateral deflections along the second direction under the dielectrophoresis force of one or more electrode areas 201, and flow out from the corresponding bypass flow channel 300 or the liquid outlet 103. That is, under the cumulative dielectrophoresis force of one or more electrode areas 201, each target biological particle in the main flow channel 100 will flow out from the preset bypass flow channel 300 or the liquid outlet 103, and the sum of the number of bypass flow channels 300 and the number of liquid outlets 103 corresponds to the number of target biological particle categories that need to be screened in the main flow channel 100.

[0058] The adjustment of the initial position of the particles is achieved by introducing a passive bypass flow field shunt mechanism, that is, a plurality of bypass flow channels 300 are sequentially communicated with the main flow channel 100 along the first direction, each bypass flow channel 300 will take away part of the flow of the main flow channel 100 (bypass shunt) to form a plurality of bypass flow fields, and make the target particles to be separated at present flow out after entering the bypass flow channel 300 through the bypass flow field, and make the next level target particles deflect to a direction closer to the first electrode 211 after passing through the current bypass flow channel 300, so that the target particles of this level are in a region with a larger electric field gradient, and will also be subjected to a stronger dielectrophoresis force.

[0059] When the sample liquid in the main flow channel 100 flows through the first level electrode area (here, the first level electrode area is the electrode area closest to the inlet channel of the main flow channel 100, and the next level electrode area will be the electrode area arranged along the first direction), the dielectrophoresis force will guide the largest particles to deflect and separate through the nearest bypass flow channel 300. After the main flow channel 100 is shunted for the first time through the bypass flow channel 300, the sample liquid will be closer to the first electrode 211, and when entering the next level electrode area, the largest particles in the remaining particles will deflect to the side close to the first electrode 211 due to the stronger dielectrophoresis force compared to the previous electric field. By repeating this process, a plurality of biological particles with a wide size distribution can be effectively separated in a microfluidic chip with a relatively simple configuration.

[0060] That is to say, the main flow channel 100 corresponding to each electrode area 201 can be divided into a bypass flow field close to the first electrode 211 and a main flow field away from the first electrode 211 in the second direction. After the target biological particles enter the corresponding electrode area 201, the initial positions thereof are generally in the main flow field, and then the target biological particles are offset from the main flow field to the bypass flow field under the action of dielectrophoresis force, and then flow out through the bypass flow channel 300. Here, the target particles and the electrode area 201 corresponding thereto correspond to each other. After the non-target particles enter the electrode area 201 corresponding to the target particles, although the non-target particles are also offset from the main flow field to the direction close to the bypass flow field under the action of dielectrophoresis force, the non-target particles will not enter the bypass flow field, and will be offset to the direction closer to the bypass flow field after passing through the current bypass flow channel 300, and then will be offset to the bypass flow field and flow out through the bypass flow channel 300 corresponding to the bypass flow field after entering the electrode area 201 corresponding thereto, or will directly flow out from the liquid outlet 103 corresponding to the outlet channel under the action of the cumulative dielectrophoresis force of the plurality of electrode areas 201.

[0061] The one or more electrode areas 201 in the separation process of the plurality of biological particles with different sizes have substantially the same electric field configuration, and it is not necessary to separately configure the electric field condition for the particles in a specific size range as in the prior art, which greatly simplifies the structure of the electrode flow channel 200 and reduces the complexity of chip setting, preparation and electric field operation.

[0062] In addition, by arranging two liquid outlets 103 at the outlet channel, the two liquid outlets 103 are communicated with the end of the main flow channel 100 in the "Y" shape in the second direction. By communicating the end of the main flow channel 100 with the two liquid outlets 103 in the "Y" shape in the second direction, the entire microfluidic chip can realize on-chip separation of N+2 different particles under the same electric field configuration. Wherein, N is the number of bypass flow channels 300 and electric field areas. The purpose of this structural design is to utilize the synergistic effect of dielectrophoresis and flow field diversion, which not only enhances the particle separation capacity of the microfluidic chip, but also improves the particle separation efficiency and accuracy in the processing of complex biological samples, and provides effective technical support for efficient biological analysis.

[0063] It should be noted that the number of liquid outlets 103 connected to the end of the main flow channel 100 is not specifically limited, and can be one liquid outlet 103 or a plurality of liquid outlets 103. However, it is more beneficial to enhance the particle separation capacity of the microfluidic chip by communicating the end of the main flow channel 100 with the two liquid outlets 103 in the "Y" shape in the second direction, without affecting the particle separation efficiency and accuracy in the complex biological samples.

[0064] Meanwhile, the design of different electrode zones 201 in the present application is also compatible with different first sub-electrode structures, thereby making different electrode zones 201 have different electric field configurations. As an example, by adjusting the flow channel length and flow channel width of each sub-electrode flow channel, the flow rate of liquid metal in each sub-electrode flow channel can be correspondingly adjusted, and thereby the input voltage of each first sub-electrode can be adjusted. The input voltage of each first sub-electrode directly affects the size of the non-uniform electric field generated by each electrode zone 201, thereby affecting the offset distance of different target particles in the second direction. That is, under the condition that the total input voltage of the first electrode 211 is stable, adjusting the flow channel length and flow channel width of each sub-electrode flow channel can adjust the size of the non-uniform electric field generated by the corresponding electrode zone 201.

[0065] Based on the above, it can be known that the key to realizing multi-particle separation lies in three main parameters, namely: the bypass flow field width w b , the initial position d0 of particles entering the electrode zone 201, and the cumulative dielectrophoresis offset distance d DEP of particles in the second direction after passing through the electrode zone 201 in the first direction.

[0066] Specifically, w b determines the degree of change of the particle flow path after each shunting of the bypass flow channel 300, and the width must be accurately designed to ensure that the separated particles smoothly exit from the bypass, and the remaining target particles can be correctly deflected to the predetermined initial position when reaching the next electrode zone. To ensure that the target particles can enter the preset bypass flow channel 300, a constraint condition needs to be met in the design, that is, the bypass flow field width w b needs to be greater than the radius of the particle to be separated, that is:

[0067] w b > r (1)

[0068] Secondly, d0 is the initial position of the particle when entering a certain level of electric field zone, which is directly related to whether the particle can enter the effective range of the non-uniform electric field. That is, when the initial position d0 of the particle is within the effective range of the non-uniform electric field, the particle will be offset to the range of bypass shunting under the action of the dielectrophoresis force of the corresponding electric field zone, and flow out from the corresponding bypass flow channel 300.

[0069] Finally, d DEP is the key to realizing precise separation, and needs to be adjusted according to the dielectric properties and size of the target particles to adjust the configuration of the first electrode array and the second electrode array, so as to ensure that the particles to be separated can obtain sufficient offset in the second direction before reaching the preset bypass flow channel 300. According to the flow line distribution rule, the condition for particle bypass separation is:

[0070] d0-d DEP <wb , (2)

[0071] Therefore, for each bypass flow channel 300 designed, the efficient separation of multiple particles from the sample liquid in the same microfluidic chip can be achieved by precisely controlling the three parameters.

[0072] In some embodiments, as shown in FIG. 8(a), each bypass flow channel 300 includes a straight flow output portion, a serpentine flow channel portion, and a straight flow connection portion, the straight flow connection portion being connected to the main flow channel 100, and the straight flow output portion being connected to a corresponding outlet. Figure 4 (a) shows that each bypass flow channel 300 includes a straight flow output portion, a serpentine flow channel portion, and a straight flow connection portion, the straight flow connection portion being connected to the main flow channel 100, and the straight flow output portion being connected to a corresponding outlet.

[0073] As can be appreciated by those skilled in the art, the width w b of the bypass flow field is related to the flow resistance of the corresponding segment of the main flow channel 100, and this relationship is determined by the geometry of the corresponding bypass flow channel 300. By providing a serpentine flow channel portion between the straight flow connection portion and the straight flow output portion, the overall flow channel length and the flow channel width of the bypass flow channel 300 can be controlled within a small range, so that the flow channel length and the flow channel width of each bypass flow channel 300 can meet the requirements of the flow channel length and the flow channel width of the segment of the main flow channel 100 within the corresponding electrode region 201. That is, by controlling the flow channel length and the flow channel width of the segment of the main flow channel 100 within the electrode region 201, the width w b of the bypass flow field within the corresponding range of the electrode region 201, and the flow channel length and the flow channel width of the corresponding bypass flow channel 300 can be determined.

[0074] In order to accurately design the actual structure of the bypass flow channel 300 to meet the requirements of separating particles of multiple sizes, it is necessary to accurately model the fluid flow in the structures of multiple bypass flow channels.

[0075] Based on the following assumptions: the flow in the main flow channel is a stable laminar flow driven by pressure; the fluid is an incompressible Newtonian fluid; and the fluid velocity distribution in the flow channel is a two-dimensional parabola, a fluid dynamics model can be used to describe the distribution of the flow velocity along the Y-axis in the main flow channel:

[0076]

[0077] where μ is the fluid viscosity, is the pressure gradient between the two ends of the channel, w t is the width of the main flow channel and 0≤y≤w t . For convenience of modeling, the bypass flow channel closest to the end of the main flow channel is set as the first bypass flow channel, and based on the continuity equation of mass conservation, the shunt width of the nth bypass flow channel depends on the ratio of the shunt flow rate at the bypass flow channel to the total flow rate before shunting, i.e.

[0078]

[0079] wherein Q tn is the total flow rate of the main channel before the diversion of the nth bypass flow channel, Q bn is the diversion flow rate of the nth bypass flow channel, v tn and w tn are the average fluid flow rate and the channel width of the main channel 100 before the diversion of the nth bypass flow channel, v bn and w bn are the average fluid flow rate and the width of the part to be diverted of the main channel 100 before the diversion of the nth bypass flow channel, and h is the channel height. Assuming that the width of the main channel 100 is constant, i.e., w t , equation (3) is substituted into equation (2) to obtain:

[0080]

[0081] By integrating the velocity profile along the Y-axis, the relationship between V and V is obtained as follows:

[0082]

[0083] Therefore, the diversion width of each bypass flow channel is determined by the diversion flow rate ratio, and the diversion flow rate ratio is determined by the flow resistance relationship between each bypass flow channel and the corresponding main channel section. In order to accurately analyze the flow resistance relationship in the flow channel structure, an equivalent circuit model of the flow channel structure is constructed, as shown in Figure 4 (b). Assuming that the geometric structures of the different bypass flow channels diverting the main channel section are the same, it is assumed that the corresponding flow resistances are also equal, i.e., R. The flow resistance of the nth bypass flow channel can be represented as B n R, wherein B n is a to-be-calculated coefficient. According to the formula of parallel resistances, the flow resistance of the flow channel network after the nth bypass flow channel can be equivalently represented by the following recursive formula:

[0084] R1=R

[0085]

[0086] Taking the diversion position of the nth bypass flow channel as a node, the following relationship can be obtained according to Kirchhoff's voltage law:

[0087]

[0088] Substituting equation (6) and equation (7) into equation (8), the bypass flow channel flow resistance coefficient B bn is obtained when the diversion width of the nth bypass flow channel is w n .

[0089] The bypass flow channel width is set as the main channel width For example (i.e. ), the flow resistance coefficient of each bypass can be calculated as shown in Figure 5 . The main flow channel is taken as an example of a rectangular cross-section, and the flow resistance of the main flow channel 100 can be calculated by the following formula:

[0090]

[0091] where f Re is the laminar flow friction constant, l is the flow channel length, D h is the hydraulic diameter of the channel, and for a rectangular channel, A is the cross-sectional area of the channel. According to formulas (6), (8) and (9), only the flow channel width and the flow channel length of the main flow channel between the bypass flow channels need to be determined to determine the flow channel width and the flow channel length of the bypass flow channels that meet the width of the bypass flow field.

[0092] From the outside, the width w b of the bypass flow field can also be adjusted by changing the height of the bypass flow channel or introducing specific structures (such as contraction-expansion micro-flow channel structures, micro-column arrays, etc.) in the bypass flow channel, which is not limited in the present application.

[0093] In some embodiments, the inlet channel in communication with the first end of the main flow channel 100 can include one sample inlet 101 and two sheath inlets 102, the two sheath inlets 102 are in communication with the first end of the main flow channel 100 in the second direction in the form of a "Y", and the sample inlet 101 is between the two sheath inlets 102 and in communication with one end of the main flow channel 100 in the form of a "Y" with the two sheath inlets 102, so that the sample liquid and the focused sheath liquid can be injected into the main flow channel 100 at the same time after converging in the inlet channel. By adjusting the flow rate Q of the focused sheath liquid of the two sheath inlets, the focusing position control of the target biological particles in the sample liquid input by the sample inlet 101 can be realized.

[0094] Specifically, under the assumption of simplifying analysis, as shown in Figure 6 , assuming that the target biological particles in the sample liquid are all at the middle position of the sample liquid, the focusing position of the particles can be represented as:

[0095]

[0096] where w1, w2 and w3 are the widths of the fluids input by the upper sheath inlet, the sample inlet and the lower sheath inlet in the main channel, respectively, and w t is the total width of the main flow channel. And w1 and w3 have the following relationship with the input flow rate of the corresponding inlet, i.e.

[0097]

[0098] where Q1, Q3 and Q t are the fluid flow rates of the upper sheath inlet, lower sheath inlet and total input of the three inlets, respectively. In combination with equations (10)-(12), the focusing position of the target biological particle in the main flow channel, i.e., the initial position d0, can be calculated as long as the input flow rates of the three inlets are given.

[0099] For the initial position of the particle entering the remaining electrode regions, the initial focusing position of the target biological particle in the previous electrode region can be subtracted by the corresponding bypass flow field width, which will not be described in detail here.

[0100] Figure 7 The schematic diagrams of the uniform sphere model and the single-shell sphere model of the embodiments of the present application are shown, and the principles of dielectrophoresis are described below in combination with Figure 7

[0101] The target particle in the sample solution is usually described based on the uniform sphere model, which assumes that the sphere has dielectric properties as shown in Figure 7 (a). The time-averaged dielectrophoretic force experienced by a uniformly spherical particle suspended in a medium can be expressed as follows:

[0102]

[0103] where R is the radius of the particle, ε m is the dielectric constant of the medium, E is the electric field strength, Re[·] represents the real part of a complex variable, K CM is the Clausius-Mossotti factor, which can be expressed as follows:

[0104]

[0105] where, and are the complex permittivities of the particle and the medium as a function of frequency, and the complex permittivities of the polarizable particle and the medium are obtained from and , where σ p and σ m represent the conductivities of the particle and the medium, respectively. j is the imaginary unit, and ω is the angular frequency of the alternating current signal.

[0106] Most biological cells, however, can be regarded as being composed of a membrane forming a vesicular structure, and their dielectric properties cannot be simplified by the above-mentioned uniform sphere model. The characteristics of these particles can be characterized by the single-shell sphere model, as shown in Figure 7 (b). The K CM of the single-shell sphere particle can be expressed as follows:

[0107] ​

[0108] where εeffis the effective complex permittivity of the particle with membrane instead of εm in equation (14) can be expressed as

[0109]

[0110] where, and are the complex permittivity of the cytoplasm and membrane, respectively, and R and d are the outer radius and membrane thickness of the single-shell dielectric model, respectively, is the area-specific membrane capacitance, is the area-specific membrane conductivity. σ cyto and σ mem are the conductivities of the cytoplasm and membrane, respectively.

[0111] As shown in Figure 8 (a), to generate the non-uniform electric field required for dielectrophoresis in the main flow channel, the first electrode array and the second column of each electrode region adopt an asymmetric electrode arrangement. After entering the electrode region, the particles are subjected to the action of dielectrophoresis force and gradually deflect in the direction with a larger electric field gradient, i.e., the positive electrode direction. As an example, the initial position d0 in the simulation parameters is set to 31 μm, and the electric field voltage is set to 22 V. This deflection process occurs continuously, and as the particles gradually pass through the electrodes, their deflection distance gradually increases. The shift of the final position of the particle along the y-axis compared with its initial position is represented as the cumulative dielectrophoresis deflection distance d DEP , which is a key parameter for evaluating the effect of electrode design and optimizing the electric field configuration.

[0112] Figure 8 (b) shows the deflection distance of the particles in the second direction under different voltage conditions, indicating that the final deflection distance of the particles increases with the increase of voltage in the range of 5-20 V. Specifically, when the electric field voltage increases, the dielectrophoresis force acting on the particles increases, resulting in an increase in the movement distance of the particles along the electric field gradient direction. When the voltage is 5 V, the deflection distance of the particles is relatively small, only about 0.4 μm, while when the voltage increases to 20 V, the deflection distance increases significantly, about 12 μm, indicating that the electric field strength has a significant impact on the dielectrophoresis deflection effect

[0113] Figure 8 (c) shows the cumulative dielectrophoresis deflection distance d DEP under different initial positions d0 and different electric field voltage settings. As can be seen from the figure, with the increase of d0, the d DEPsignificantly. For example, at a voltage of 10 V, the deflection distance of the particle decreases sharply from 7 μm to 2.4 μm when d0 increases from 20 μm to 25 μm, which indicates that the initial position of the particle has a significant effect on the cumulative dielectrophoretic deflection. To satisfy the condition of the particle bypass separation proposed in the present application, the bypass flow field width is set to be 10 μm as an example. In the figure, the dashed line represents that the initial position of the particle minus the cumulative dielectrophoretic deflection distance is just equal to the bypass flow field width. The shaded area above the dashed line represents the parameter settings that satisfy the separation condition, and if the initial position of the particle and d DEP fall within the shaded area, effective separation of the target particles can be achieved.

[0114] According to a second aspect of the present application, a microfluidic chip-based biological particle separation device is provided, comprising the microfluidic chip according to any one of the above embodiments.

[0115] In a specific embodiment, as shown in Figure 9 , the sample fluid is an artificially mixed sample including HeLa cells with a diameter of 15-20 μm, white blood cells (WBCs) with a diameter of 10-15 μm, red blood cells (RBCs) with a diameter of 6-8 μm, and PS microspheres with a diameter of 0.5 μm. These particles with a size gradient distribution are selected to simulate the main particle distribution in blood and verify the separation capability of the chip structure proposed in the present application based on particle size. In addition, in order to verify that the chip also has the separation capability based on the dielectric properties of the particles, 10-μm-diameter PS microspheres are added to the experimental sample to simulate the immunomagnetic beads commonly used in biomedical detection research. This multi-level particle selection not only reflects the adaptability of the microfluidic chip provided in the present application to complex biological sample processing, but also demonstrates its high efficiency under the dual separation mechanism of size and dielectric properties. At the same time, HeLa cells are stained with fluorescent dye Calcein AM, and WBCs are stained with Hoechst-33342 to facilitate particle tracking.

[0116] As shown in Figure 9 (a-b), when an electric field with a voltage of 40 Vpp and a frequency of 1 MHz is applied, the multi-particle separation effect of the microfluidic chip is clearly visible. The largest HeLa cells in the mixed sample are first successfully separated at outlet one (01), followed by WBCs, which are separated from outlet two (02). Smaller RBCs are separated from outlet three (03), and even smaller 0.5-μm-diameter PS microspheres are mainly separated through outlet four (04). In addition, 10-μm-diameter PS microspheres similar in size to white blood cells are pushed away from the main sample flow under the action of nDEP force at this electric field setting, and are finally separated through outlet five (05). Figure 9(c-d) The change of sample purity before and after separation proves the ability of the chip structure proposed in the application to separate particles of different sizes in complex samples.

[0117] In summary, the application provides a microfluidic chip and a biological particle separation device based on the microfluidic chip, which comprises a main flow channel 100, a plurality of bypass flow channels and an electrode flow channel 200. The electrode flow channel 200 is arranged on both sides of the side wall of the main flow channel 100 and forms a plurality of electrode zones 201 along the extension direction of the main flow channel 100, thereby providing a non-uniform electric field for the main flow channel 100. The plurality of bypass flow channels are sequentially communicated with the main flow channel 100 along the extension direction of the main flow channel 100, so that the electrode zones 201 and the bypass flow channels can be arranged alternately along the extension direction of the main flow channel 100. The application combines the passive flow field shunting mechanism with the active dielectrophoresis accumulation and deflection effect, so that different target biological particles in the sample liquid injected into the main flow channel 100 can produce different lateral deflections in the normal direction of the flow direction of the sample liquid under the action of the dielectrophoresis force of one or more electrode zones 201, so that the target biological particles can flow out through the corresponding bypass flow channel or liquid outlet 103, thereby realizing efficient separation of biological particles of different sizes and types.

[0118] The microfluidic chip and the biological particle separation device based on the microfluidic chip provided by the application combine the passive flow field shunting mechanism with the active electric field separation mechanism, thereby reducing the manufacturing cost and design complexity of the chip or device, and simplifying the configuration and regulation requirements of the physical field.

[0119] The microfluidic chip and the biological particle separation device based on the microfluidic chip provided by the application are also compatible with multiple separation mechanisms, so that they have higher versatility when processing multiple different types of particles. Whether the particles are not easy to mark or the particles are of different sizes and physical properties, efficient separation can be realized in the same device.

[0120] In the foregoing embodiment description, the description of the terms "one embodiment", "some embodiments", "an example", "a specific example", or "some examples" and the like means that the specific features, structures, materials or characteristics described in connection with the embodiment or example are included in at least one embodiment or example of the application. In the specification, the illustrative description of the above terms does not necessarily refer to the same embodiment or example. Moreover, the specific features, structures, materials or characteristics described can be combined in any appropriate manner in any one or more embodiments or examples. In addition, the person skilled in the art can combine and combine the different embodiments or examples described in the specification and the features of the different embodiments or examples without contradiction.

[0121] Furthermore, the terms "first", "second", "third", "fourth", "fifth", "sixth", "seventh" and "eighth" are used herein for descriptive purposes only and are not to be construed as indicating or implying relative importance or a significant nature of so described technical features. It is to be expressly understood that a description with "first", "second", "third", "fourth", "fifth", "sixth", "seventh" and "eighth" features can implicitly or explicitly include at least one of such features. In the description of the present application, the meaning of "a plurality" is at least two, for example, two, three or more, unless otherwise specifically defined.

Claims

1. A microfluidic chip, characterized by, The microfluidic chip comprises a main flow channel, a plurality of bypass flow channels and an electrode flow channel; wherein One end of the main flow channel is connected with a sample inlet and a plurality of sheath liquid inlets, and the opposite end is connected with a plurality of liquid outlets; sample liquid and sheath liquid are injected into the main flow channel through the sample inlet and the sheath liquid inlet respectively; The electrode flow channel is arranged on both sides of the side wall of the main flow channel in the second direction, and forms a plurality of electrode zones in the first direction, each of which provides an uneven electric field for the main flow channel at the corresponding position in the second direction, the first direction being the extension direction of the main flow channel, and the second direction being perpendicular to the extension direction of the main flow channel; A plurality of bypass flow channels are sequentially connected with the main flow channel in the first direction, and each bypass flow channel corresponds to each electrode zone one by one, and the electrode zone and the bypass flow channel are alternately spaced in the first direction, so that different target biological particles in the sample liquid injected into the main flow channel can produce different lateral deviations in the second direction under the action of dielectrophoresis force of one or more electrode zones, and flow out from the corresponding bypass flow channel or liquid outlet; The electrode flow channel comprises a first electrode flow channel and a second electrode flow channel, and the first electrode flow channel and the second electrode flow channel are respectively filled with liquid metal to form the first electrode and the second electrode; different target biological particles in the main flow channel move to the side close to the first electrode in the second direction under the action of dielectrophoresis force of one or more electrode zones; The first electrode comprises a plurality of first sub-electrodes, and a plurality of first sub-electrodes and the second electrode correspond to each other to form a plurality of electrode zones spaced from each other; wherein the first electrode is a positive electrode, and the second electrode is a negative electrode; the main flow channel corresponding to each electrode zone comprises a bypass flow field close to the first electrode and a main flow field away from the first electrode in the second direction, and when the target biological particles enter the bypass flow field from the main flow field under the action of dielectrophoresis force of the electrode zone and satisfy the following relationship, the target biological particles will flow out from the bypass flow channel corresponding to the electrode zone, that is: d 0 -d DEP <w b ， wherein, d 0 is an initial position of the target biological particle along the second direction when entering the electrode region, d DEP is an offset of the target biological particle along the second direction under the action of the dielectrophoresis force in the electrode region, w b is a width of the bypass flow field along the second direction.

2. The microfluidic chip of claim 1, wherein, One end of the main flow channel is connected with at least one sample inlet and at least two sheath liquid inlets, and the other end of the main flow channel is connected with at least one liquid outlet.

3. The microfluidic chip of claim 1, wherein, Each bypass flow channel comprises a direct current output part, a serpentine flow channel part and a direct current connection part, the direct current connection part connects the serpentine flow channel part with the main flow channel, and the direct current output part connects the serpentine flow channel part with a corresponding outlet.

4. The microfluidic chip of claim 1, wherein, The radius of the target biological particle is smaller than the width of the corresponding bypass flow field in the second direction.

5. The microfluidic chip of claim 1, wherein, The lengths and widths of the main flow channels corresponding to different electrode zones are not the same, and the lengths and widths of each bypass flow channel correspond to the lengths and widths of the main flow channel corresponding to the electrode zone.

6. A microfluidic chip-based biological particle separation device, characterized by, The microfluidic chip comprises the microfluidic chip according to any one of claims 1-5.

Images