Culture device and method for automated expansion of cells

Abstract

This invention relates to the field of cell culture and bioengineering equipment, specifically to a cell culture device and method for automated cell expansion; it includes a rack, a culture chamber, a transducer array, a perfusion pump, and a controller; the system combines the resonant frequency drift of the transducer array with the torque fluctuation decay time of the perfusion pump stepper motor to calculate cell confluence and local shear stress; its core is to achieve non-contact data monitoring and calculation without introducing invasive probes and ensuring a sterile system; this invention abandons traditional invasive detection and uses non-contact calculation to provide accurate data for adjusting system operating parameters, effectively overcoming the technical defects of existing equipment that cannot obtain local rheological properties and the true growth state of cells in real time.

Description

Technical Field

[0001] This invention relates to the field of cell culture and bioengineering equipment, specifically to a cell culture device and method for automated cell expansion. Background Technology

[0002] In the current research environment of bioengineering, high-density cell expansion generally requires the relative movement of microcarriers and culture medium for material exchange. To meet the high mass transfer requirements during rapid cell proliferation, existing cell expansion equipment typically enhances mass transfer by simply increasing the perfusion rate or strengthening the stirring intensity. Although this approach improves the mass transfer efficiency within the system to some extent, simply enhancing the action of a single fluid can easily generate local high-shear zones on the surface of the microcarrier. These local high-shear forces directly impact the interface, causing attached cells to detach or suffer mechanical damage. Furthermore, existing equipment struggles to acquire real-time data on cell confluence and local rheological properties without contact or the introduction of invasive probes, resulting in a lack of accurate basis and dynamic coordination mechanism for adjusting system operating parameters.

[0003] Therefore, how to effectively resolve the conflict between mass transfer requirements and shear damage during cell expansion, and how to achieve adaptive dynamic scheduling of force field parameters while ensuring a sterile system, has become an urgent technical problem to be solved. Summary of the Invention

[0004] To solve the above-mentioned technical problems, the present invention provides a cell culture device and method for automatic cell expansion. Specifically, the technical solution of the present invention is as follows: A cell culture device for automated cell expansion, comprising: The rack, including the lower partition; The culture chamber is fixed in the center of the frame by a flange. The culture chamber includes a cylindrical quartz glass tube and an inner liner, wherein the inner liner is coaxially sleeved inside the cylindrical quartz glass tube. A transducer array is fixed to the outer wall of the cylindrical quartz glass tube; A perfusion pump is fixed to the lower partition of the frame. The perfusion pump includes a stepper motor, and the output end of the perfusion pump is connected to the bottom of the culture chamber. The controller is connected to control the operation of the transducer array and the perfusion pump.

[0005] Furthermore, the outer wall of the inner liner is interference-fitted with the inner wall of the cylindrical quartz glass tube, and the inner wall of the inner liner is arrayed with inverted conical blind holes, wherein the openings of the inverted conical blind holes face the center of the culture chamber.

[0006] Furthermore, the inner liner is made of polyetheretherketone material, the cone angle of the inverted conical blind hole is 50° to 70°, and a polyurethane elastic film is hot-melt welded to the open end of the inverted conical blind hole. The thickness of the polyurethane elastic film is 0.05mm to 0.20mm. The polyurethane elastic film seals the interior of the inverted conical blind hole to form an inverted conical air chamber with initial air pressure.

[0007] Furthermore, the transducer array includes piezoelectric ceramic transducers, which are fixedly connected to the outer wall of the cylindrical quartz glass tube in orthogonal directions on the same horizontal plane by conductive silver paste, wherein the piezoelectric ceramic transducers form orthogonal ultrasonic emission pairs in pairs.

[0008] Furthermore, the culture chamber also includes a bottom end cap and a top end cap. The bottom end cap is machined with a liquid inlet channel, which is connected to the output end of the perfusion pump via a silicone hose. The perfusion pump uses a stepper motor to directly drive the peristaltic pump head. The top end cap of the culture chamber is machined with a liquid outlet channel.

[0009] A method for controlling the automatic proliferation of cells in a culture, comprising: S1. Control the perfusion pump to continuously input pulsed fluid into the culture chamber, and control the transducer array to establish an orthogonal ultrasonic standing wave field in the culture chamber; S2. Obtain the current resonant frequency of the transducer array and the torque fluctuation decay time of the stepper motor through the controller; S3. Calculate the cell confluence degree based on the drift of the resonant frequency of the transducer array. S4. Calculate the local shear stress based on the torque fluctuation decay time of the stepper motor of the irrigation pump; S5. Dynamically schedule the force field parameters of the transducer array and the perfusion pump based on the calculated cell confluence and local shear stress.

[0010] Furthermore, methods for calculating cell confluence include: Obtain the reference resonant frequency in the initial cell-free state; calculate the frequency drift by subtracting the currently measured resonant frequency from the reference resonant frequency in the initial cell-free state. The cell confluence is obtained by multiplying the frequency drift by a compensation coefficient related to the sound speed of the medium.

[0011] Furthermore, methods for calculating local shear stress include: Multiplying the torque fluctuation decay time by the flow resistance coefficient related to the pipe cross-sectional area yields the fluid dynamic viscosity. The average flow velocity in the pipeline is calculated based on the current rotational speed of the irrigation pump; The calculated average flow velocity in the pipeline is multiplied by the calculated dynamic viscosity of the fluid and divided by the pipeline characteristic scale factor to calculate the local shear stress inside the culture chamber; wherein, the pipeline characteristic scale factor is an equivalent length parameter characterizing the hydraulic characteristics of the pipeline cross section, which can be obtained through calibration experiments with blank culture medium.

[0012] Furthermore, the inner wall of the inner liner of the culture device is machined with an inverted conical blind hole, the opening end of which is sealed with a polyurethane elastic film to form an inverted conical air chamber; the step of dynamically scheduling the force field parameters of the transducer array and the perfusion pump according to the calculated cell confluence degree and the local shear stress includes: When the calculated cell confluence is less than a preset confluence threshold, the transducer array is controlled to output basic maintenance power, and the stepper motor of the perfusion pump is controlled to run at a high pulse width duty cycle to output a high pulse flow rate, so that the high pulse fluid squeezes the polyurethane elastic film on the inner liner into the inverted conical air chamber.

[0013] Furthermore, a safety error envelope is set, which is a pre-defined local shear stress allowable range that meets the target cell viability requirements; the step of dynamically scheduling the force field parameters of the transducer array and the perfusion pump based on the calculated cell confluence and local shear stress also includes: When the calculated cell confluence is greater than or equal to the preset confluence threshold, the transducer array is controlled to increase the transmission power, while the pulse width duty cycle of the perfusion pump is shortened and the pulse flow rate is reduced. When the calculated local shear stress reaches the upper limit of the safety error envelope, the pulse width duty cycle is reduced or the transmit power of the transducer array is increased to keep the local shear stress within the safety error envelope.

[0014] The present invention has the following beneficial effects: 1. This invention obtains the resonant frequency drift of the transducer array and the torque fluctuation decay time of the perfusion pump stepper motor through a controller. Without introducing invasive probes and ensuring a sterile system, it realizes non-contact calculation of cell confluence and local shear stress. This method provides accurate data for adjusting system operating parameters and effectively overcomes the technical defects of existing equipment that make it difficult to obtain local rheological properties and the actual cell growth state in real time. 2. This invention dynamically adjusts the force field parameters based on the calculated cell confluence and local shear stress. When the confluence is low, a high-pulse fluid is used to compress the polyurethane elastic film on the inner liner into the inverted conical air chamber to absorb the pulse impact. When the confluence is high, the transducer array emission power is increased and the perfusion pump pulse width duty cycle is shortened to reduce the flow rate. This dynamic coordination mechanism successfully resolves the conflict between high mass transfer requirements and shear damage, ensuring that the shear stress on the cells remains within the safe error envelope. Attached Figure Description

[0015] Figure 1 This is a schematic diagram of the overall external structure of the device; Figure 2 This is a schematic diagram of the lower partition and its connection structure; Figure 3 This is a schematic diagram of the culture chamber structure; Figure 4 This is a flowchart of the method of the present invention.

[0016] In the diagram: 1. Frame; 2. Lower partition; 3. Culture chamber; 4. Flange; 5. Cylindrical quartz glass tube; 6. Inner liner; 8. Transducer array; 9. Perfusion pump; 10. Controller; 11. Inverted conical blind hole; 12. Polyurethane elastic film; 13. Inverted conical gas chamber; 14. Piezoelectric ceramic transducer; 15. Conductive silver paste; 16. Bottom end cap; 17. Inlet channel; 18. Silicone tubing; 19. Stepper motor; 20. Peristaltic pump head; 21. Top end cap; 22. Outlet channel. Detailed Implementation

[0017] The technical solutions in the embodiments of this application will be clearly and completely described below with reference to the accompanying drawings. Obviously, the described embodiments are only some embodiments of this application, not all embodiments. All other embodiments obtained by those skilled in the art based on the embodiments of this application without creative effort are within the scope of protection of this application. Example

[0018] Combination Figure 1 A cell culture device for automated cell proliferation, comprising: Frame 1, including lower partition 2; The culture chamber 3 is fixed to the center of the frame 1 by a flange 4. The culture chamber 3 includes a cylindrical quartz glass tube 5 and an inner liner 6, wherein the inner liner 6 is coaxially sleeved inside the cylindrical quartz glass tube 5. A transducer array 8 is fixed to the outer wall of a cylindrical quartz glass tube 5; a perfusion pump 9 is fixed to the lower partition 2 of the frame 1, and the perfusion pump 9 includes a stepper motor 19, wherein the output end of the perfusion pump 9 is connected to the bottom of the culture chamber 3. Controller 10 is connected to control the operation of transducer array 8 and perfusion pump 9; Existing high-density cell expansion equipment typically enhances mass transfer by simply increasing the perfusion rate or stirring intensity. This method can easily create local high-shear zones on the surface of the microcarrier, leading to cell detachment or mechanical damage. To address this issue, a frame 1 is set up as the foundation for the entire machine. The frame 1 is made of a corrosion-resistant metal frame, preferably welded from 304 stainless steel or 316L stainless steel. The lower partition 2 is used to support the perfusion pump 9 and the controller 10, so that the power components and the culture chamber 3 are arranged in layers vertically to reduce the direct transmission of vibration to the culture chamber 3. The culture chamber 3 is fixed to the center of the frame 1 by a flange 4. The flange 4 connection allows the culture chamber 3 to be disassembled and easily replaced after sterilization. The culture chamber 3 includes a cylindrical quartz glass tube 5 and an inner liner 6. The cylindrical quartz glass tube 5 serves as an ultrasonic transmission medium and a sterile barrier, with its axis arranged vertically. The inner liner 6 is coaxially arranged inside the cylindrical quartz glass tube 5 to form a microstructured compliant interface. The internal space of the culture chamber 3 is used to accommodate microcarriers with attached cells. These microcarriers are aggregated and distributed on the inside of the cylindrical quartz glass tube 5 under the action of gravity and perfusion, thereby forming a microcarrier bed area through which the culture medium flows. The transducer array 8 is fixed to the outer wall of the cylindrical quartz glass tube 5 and is used to input ultrasonic energy into the culture chamber 3 and establish a predetermined sound field; the perfusion pump 9 is fixed on the lower partition 2 and its output end is connected to the bottom of the culture chamber 3 through a sterile connecting pipe, so that the culture medium passes through the microcarrier bed area from bottom to top. The controller 10 is electrically connected to the transducer array 8 and the irrigation pump 9 respectively. The controller 10 is preferably composed of an industrial microprocessor, a motor drive module, a power amplifier module and a data acquisition module. It is used to issue pumping pulse commands, adjust the transducer transmission power, read the resonant frequency and motor operating status, and schedule the equipment operating parameters according to the set algorithm. Through the above structural combination, the culture device concentrates the fluid and acoustic effects within the same culture chamber 3, providing a structural basis for the synergistic control of high mass transfer and low shear.

[0019] Combination Figure 3 The outer wall of the inner liner 6 is press-fitted with the inner wall of the cylindrical quartz glass tube 5. The inner wall of the inner liner 6 is arrayed with inverted conical blind holes 11, wherein the opening of the inverted conical blind holes 11 faces the center of the culture chamber 3. To reduce the disturbance of transient pressure waves caused by pulsed perfusion on the stability of standing waves inside the culture chamber 3, the outer wall of the inner liner 6 and the inner wall of the cylindrical quartz glass tube 5 are set to an interference fit. An interference fit means that the outer diameter of the inner liner 6 is larger than the inner diameter of the cylindrical quartz glass tube 5 by a preset interference amount, so that a reliable fit is formed after assembly. The preferred interference amount is 0.02 mm to 0.08 mm. Within this range, the continuity of acoustic vibration transmission can be guaranteed, and the cracking of the cylindrical quartz glass tube 5 due to excessive assembly stress can also be avoided. The inner wall of the inner liner 6 is arrayed with inverted conical blind holes 11. The array processing refers to forming multiple repeating units at a set interval along the circumferential and axial directions. Preferably, one column is set every 90° in the circumferential direction, and the axial spacing is 3mm to 8mm. The opening of the inverted conical blind hole 11 faces the center of the culture chamber 3, that is, the large diameter end of the blind hole faces the mainstream area of ​​the culture medium, and the small diameter end faces the inside of the inner liner 6. This arrangement allows for controllable local surface deformation when the pulsed fluid acts on the opening area. Compared with ordinary cylindrical holes, the inverted conical blind hole 11 can achieve improved compliance adjustment capability under the same hole depth conditions. This is because its cross-sectional area gradually shrinks along the depth direction, and the membrane surface deflection formed after the fluid pressure is applied is oriented and concentrated, which can stably absorb water hammer pulses. This structure also enables the ultrasonic waves to form adjustable boundary compliance conditions when passing through the interface of the inner liner 6, which is beneficial to maintaining the nodal stability of the orthogonal standing wave field in the cavity.

[0020] The inner liner 6 is made of polyetheretherketone material. The cone angle of the inverted conical blind hole 11 is 50° to 70°. The opening end of the inverted conical blind hole 11 is hot-melt welded with a polyurethane elastic film 12. The thickness of the polyurethane elastic film 12 is 0.05mm to 0.20mm. The polyurethane elastic film 12 seals the interior of the inverted conical blind hole 11 to form an inverted conical air chamber 13 with initial air pressure. The inner liner 6 is preferably made of polyetheretherketone (PEEK) material. PEEK material refers to engineering plastics that can withstand high-temperature steam sterilization, have mechanical strength that meets the system pressure requirements, and constant acoustic properties. It is not easily deformed under long-term immersion in culture medium and is suitable for maintaining the stability of blind hole array parameters. The cone angle of the inverted conical blind hole 11 is set to a preset angle, preferably 50° to 70°, more preferably 60°; by controlling the cone angle within this range, a stable correspondence can be formed between the effective volume of the blind hole, the pressure area of ​​the diaphragm, and the compression stiffness of the air chamber. The opening end of the inverted conical blind hole 11 is hot-melt welded with a polyurethane elastic film 12. The polyurethane elastic film 12 refers to a polymer film material with reversible elastic deformation capability and capable of forming a sealed connection with the edge of the polyether ether ketone opening. The thickness is preferably 0.05 mm to 0.20 mm, and more preferably 0.10 mm. The hot melt welding process can use a heating head for local heating, and maintain the temperature at 160°C to 220°C for 0.5s to 3s to obtain a continuous weld; the thin film seals the inside of the blind hole to form an inverted conical gas chamber 13 with an initial gas pressure, which is preferably 0.9 atm to 1.2 atm; The initial gas pressure is not set arbitrarily, but is determined together with the film thickness, cone angle and predetermined perfusion pulse pressure. In order to offset the gas contraction pressure drop caused by the high temperature cooling after hot melt welding, and to ensure that the gas chamber can accurately reach the initial gas pressure under normal operating conditions, the hot melt welding process is carried out in a sealed chamber with pressure regulation function. By filling the chamber with high-pressure gas with calculated temperature compensation during the welding stage, a pressure pre-embedded to match the temperature change is provided to control the compression of the gas chamber when subjected to water hammer pulses. In this embodiment, the water hammer pulse specifically refers to the transient pressure fluctuation of the fluid inside the confined space of the pipeline and culture chamber 3 when the perfusion pump 9 is driven by the periodic pulse of the stepper motor 19 to deliver viscous culture medium at a non-constant flow rate; when the pulsed fluid generates transient high pressure, the polyurethane elastic film 12 is recessed into the inverted conical air chamber 13, the volume of the air chamber decreases and the pressure energy is stored, and the film resets and releases the energy after the pressure falls back. Thus, the walls of culture chamber 3 acquire compliance characteristics that match transient pressure, which can reduce the impact of pressure peaks on the cell attachment interface and reduce ultrasound standing wave phase fluctuations.

[0021] The transducer array 8 includes piezoelectric ceramic transducers 14. The piezoelectric ceramic transducers 14 are fixedly connected to the outer wall of the cylindrical quartz glass tube 5 in orthogonal directions on the same horizontal plane by conductive silver paste 15. The piezoelectric ceramic transducers 14 form orthogonal ultrasonic emission pairs in pairs. The transducer array 8 preferably includes four piezoelectric ceramic transducers 14; the piezoelectric ceramic transducer 14 refers to an acoustic-electric conversion element that generates mechanical vibration under AC drive and couples ultrasonic waves to the cylindrical quartz glass tube 5. It is preferably a sheet or arc-shaped transducer made of lead zirconate titanate material, and the operating frequency can be set in the range of 0.5MHz to 5MHz. Four piezoelectric ceramic transducers 14 are located in orthogonal directions on the same horizontal plane, that is, one is arranged every 90° along the circumference of the cylindrical quartz glass tube 5, and their centers are located in the same cross section. They are fixedly connected to the outer wall of the cylindrical quartz glass tube 5 by conductive silver paste 15. The conductive silver paste 15 serves as both a mechanical adhesive and an electrical connection and acoustic coupling. The thickness of the paste layer is preferably controlled between 0.03 mm and 0.15 mm to reduce additional damping. Pairs of opposing piezoelectric ceramic transducers 14 form orthogonal ultrasonic emission pairs, one pair arranged along the X direction and the other pair arranged along the Y direction. The controller 10 applies a drive signal of the same frequency and a controllable phase difference to the two pairs of transducers, which can form an orthogonal ultrasonic standing wave field in the culture chamber 3. Unlike unidirectional ultrasonic excitation, this structure, with its orthogonal arrangement, can provide a two-dimensional acoustic potential distribution within the chamber cross-section, allowing the microcarrier and its surrounding culture medium to be modulated by acoustic radiation forces in multiple directions, which is beneficial for forming a more stable low-shear local region. The piezoelectric ceramic transducer 14 can also be used as a detection element for changes in acoustic load. Its impedance spectrum and resonant frequency are affected by the acoustic impedance of the medium inside the chamber. After the cell biomass increases, the change in the mechanical boundary conditions of the transducer can be reflected in the resonant peak drift, providing available raw data for subsequent cell confluence calculation.

[0022] The culture chamber 3 also includes a bottom end cap 16 and a top end cap 21. The bottom end cap 16 is machined with a liquid inlet channel 17, which is connected to the output end of the perfusion pump 9 through a silicone hose 18. The perfusion pump 9 uses a stepper motor 19 to directly drive the peristaltic pump head 20. The top end cap 21 of the culture chamber 3 is machined with a liquid outlet channel 22. Combination Figure 2 The bottom end cap 16 of the culture chamber 3 is machined with an inlet channel 17, and the top end cap 21 is machined with an outlet channel 22. Together, they define the perfusion path of the culture medium from bottom to top. The bottom inlet and top outlet arrangement allows the culture medium to pass through the area where the cells and microcarriers are located along the axis in the chamber, reducing trapped air bubbles and facilitating the discharge of metabolic products along the main flow direction. The inlet channel 17 is connected to the output end of the perfusion pump 9 via a silicone hose 18. The silicone hose 18 is preferably made of medical-grade silicone material, with an inner diameter of 1mm to 6mm, and has the characteristics of flexibility, non-toxicity, and resistance to repeated compression. The perfusion pump 9 uses a stepper motor 19 to directly drive the peristaltic pump head 20. Direct drive means that the output shaft of the stepper motor 19 is fixedly connected to the main shaft of the peristaltic pump head 20, without the intermediate belt or gear reduction structure, so as to reduce the elastic lag of the transmission. The stepper motor 19 can be driven by microstepping, with a step angle preferably of 1.8° or 0.9°. Repeatable pulse perfusion is formed by setting the pulse frequency and pulse width duty cycle. Preferably, the stepper motor 19 is a closed-loop control stepper motor 19 with an integrated high-precision position encoder, which can feed back the actual physical position of the rotor to the controller 10 in real time, thereby outputting a step loss compensation signal to characterize the magnitude of the fluid damping load. The rollers of the peristaltic pump head 20 periodically press the silicone tubing 18, so that the culture medium is delivered to the culture chamber 3 under conditions without the risk of contact contamination. Since the output torque of the stepper motor 19 changes with the viscosity of the culture medium, the direct drive structure can reflect the load disturbance more directly to the motor current, phase error or torque fluctuation signal, making it easier for the controller 10 to extract the torque fluctuation decay time. Through the fluid pathway consisting of the bottom inlet channel 17, the silicone hose 18, the stepper motor 19 driving the peristaltic pump head 20, and the top outlet channel 22, the equipment has both aseptic perfusion capability and the ability to detect fluid rheological properties indirectly from the power side. Example

[0023] Combination Figure 4 A method for controlling the automatic proliferation of cells in a culture, comprising: S1. Control the perfusion pump 9 to continuously input pulsed fluid into the culture chamber 3, and control the transducer array 8 to establish an orthogonal ultrasonic standing wave field inside the culture chamber 3. S2. The current resonant frequency of the transducer array 8 and the torque fluctuation decay time of the stepper motor 19 are obtained through the controller 10. S3. Calculate the cell confluence degree based on the drift of the resonant frequency of the transducer array 8. S4. Calculate the local shear stress based on the torque fluctuation decay time of the stepper motor 19 of the irrigation pump 9. S5. Dynamically adjust the force field parameters of transducer array 8 and perfusion pump 9 based on the calculated cell confluence and local shear stress. This control method is used to resolve the conflict between mass transfer requirements and shear damage during cell expansion; pulsed fluid refers to the non-constant flow rate culture medium output by the perfusion pump 9 according to a set rhythm; orthogonal ultrasonic standing wave field refers to the fixed spatial sound pressure distribution formed by ultrasonic waves emitted by two pairs of mutually perpendicular transducers and superimposed in the culture chamber 3. The controller 10 drives the perfusion pump 9 to continuously input pulsed fluid into the culture chamber 3. The pulse frequency is preferably 0.1Hz to 5Hz. The duration of high flow rate within a single pulse cycle is determined by the pulse width duty cycle. At the same time, the controller 10 outputs an AC drive signal to the transducer array 8 to establish an orthogonal ultrasonic standing wave field in the culture chamber 3. During cell expansion, cells attach to microcarriers and proliferate continuously, increasing the acoustic mass load of the medium inside the chamber and causing the transducer resonant frequency to drift. Cells secrete extracellular matrix and change the rheological properties of the culture medium, and the load disturbance attenuation process of stepper motor 19 changes accordingly in each pulse cycle. The controller 10 collects the impedance spectrum data of the transducer array 8, identifies the position of the resonance peak and obtains the resonant frequency drift; it collects the driving current or back electromotive force of the stepper motor 19 to calculate the torque fluctuation curve and extracts the torque fluctuation decay time; the controller 10 calculates the cell confluence based on the resonant frequency drift and calculates the local shear stress based on the torque fluctuation decay time, and then uses the two calculation results as scheduling inputs to adjust the power of the transducer array 8 and the pulse width duty cycle of the perfusion pump 9; The key point of this method is that the resonant frequency drift is not an isolated monitoring value, but is directly involved in the calculation of cell confluence. The torque fluctuation decay time is not simply a recorded value, but is directly involved in the calculation of local shear stress. Both physical quantities are included in the force field parameter scheduling process, so that the culture device can maintain the mass transfer and shear balance required for cell expansion without introducing an invasive probe.

[0024] The method for calculating cell confluence includes: obtaining the reference resonant frequency in the initial cell-free state; subtracting the currently measured resonant frequency from the reference resonant frequency in the initial cell-free state to obtain the frequency drift; and multiplying the frequency drift by a compensation coefficient related to the sound velocity of the medium to obtain the cell confluence. In this invention, cell confluence is used to characterize the degree of cell occupancy on the surface and within the pores of the microcarrier. It is not limited to the concept of coverage in planar culture, but rather represents the expansion level of cell biomass relative to the initial inoculation state. Before calculating cell confluence, the controller 10 performs a benchmark calibration after the cell-free culture medium and blank microcarriers are loaded into the culture chamber 3, and drives the transducer array 8 to perform a frequency sweep to obtain the benchmark resonant frequency in the initial cell-free state. During the amplification operation, the controller 10 repeatedly sweeps the frequency at a set time interval, which can be from 5 min to 60 min, to obtain the currently measured resonant frequency; the difference between the currently measured resonant frequency and the reference resonant frequency is used to obtain the frequency drift. Since cell attachment and proliferation increase the total acoustic mass load in the chamber, the frequency drift usually manifests as a shift towards lower frequencies, and this shift increases with the increase of cell biomass. In order to reduce the influence of different culture medium formulations, temperatures and cell types on the sound propagation velocity, the frequency drift is multiplied by a compensation coefficient related to the sound velocity of the medium to obtain the cell confluence. The compensation coefficient can be established through preliminary experiments. It is obtained by measuring multiple frequency shifts under a standard sample with known cell biomass, and then fitting a linear or piecewise linear relationship with the corresponding sound velocity measurements in the culture medium to obtain the compensation coefficient, which has the dimension of the reciprocal of frequency. Let the currently measured frequency drift be... Then the degree of cell confluence The calculation formula can be explicitly expressed as:

[0025] By fitting a linear or piecewise linear relationship with the sound velocity measurements of the corresponding culture medium, the compensation coefficient k is obtained; the sound velocity of the medium can be pre-measured by an ultrasonic velocimeter using the same batch of culture medium at a set temperature, with the temperature preferably controlled between 36℃ and 38℃; Using this method, the transducer body serves both the excitation and state recognition functions, and the acquisition of cell confluence does not require sampling and microscopic observation, thereby reducing the probability of the sterile system being destroyed. The calculation process essentially constructs a cell biomass acoustic assessment model. The purpose of this model is to accurately estimate the real-time confluence of cells on the microcarrier under sterile conditions and without introducing invasive probes. In terms of logic structure and data flow, the model receives the currently measured resonant frequency and the pre-calibrated reference resonant frequency as inputs to calculate the frequency drift; it also receives a compensation coefficient determined based on temperature and culture medium formulation as another input; and finally outputs the two sub-results as cell confluence degree through multiplication logic. In terms of the physical relationships represented, this model characterizes the physical phenomenon that when ultrasound propagates in culture chamber 3, the cells attached to the surface of the microcarrier increase the total acoustic mass load of the medium, causing the natural frequency of the resonant system to shift to a lower frequency. Since the increase in cell biomass directly changes the mechanical impedance boundary conditions of the chamber, the cell expansion level can be accurately reflected by monitoring the drift of the resonant frequency and supplementing it with sound velocity compensation. In the control logic, cell confluence is a state quantity used to determine whether the culture system is currently in a mass transfer priority phase or a low shear protection priority phase. Its source is not an instantaneous measurement at a single frequency point, but rather obtained by filtering the frequency sweep results. Specifically, the controller 10 preferably first sets a frequency sweep window near the current operating frequency, which can be 2kHz to 20kHz on both sides of the operating frequency; collects conductance, phase or impedance amplitude data at each frequency point, identifies the impedance minimum point or phase change point as candidate values ​​for the current resonant frequency; and then determines the candidate value that appears repeatedly in at least two consecutive samples as the currently measured resonant frequency, so as to reduce misjudgment caused by bubble disturbance or instantaneous flow fluctuation. The frequency drift is preferably taken as the absolute offset amplitude of the current resonant frequency minus the reference resonant frequency, and the validity is checked in combination with the drift direction; when the drift direction is detected to be inconsistent with the preset cell proliferation trend, the controller 10 does not update the cell confluence degree, but maintains the result of the previous cycle and performs a frequency scan confirmation again. The compensation coefficient not only characterizes the difference in sound velocity in the medium, but is also used to map the frequency quantity to the dimension of cell occupancy. The determination steps preferably include: Step 1, recording the frequency drift under different known cell biomass conditions; Step 2, measuring the sound velocity of the corresponding culture medium under the same temperature conditions; Step 3, establishing a corresponding table with cell biomass or occupancy ratio obtained by offline microscopic statistics as a reference; Step 4, the controller 10 calls the parameter table of the corresponding cell type and culture medium batch, and selects the compensation coefficient that is closest to the current sound velocity. Therefore, the cell confluence degree output by the process is passed to the force field parameter scheduling module and used as a direct input for comparing the preset confluence degree threshold.

[0026] The method for calculating local shear stress includes: multiplying the torque fluctuation decay time by the flow resistance coefficient related to the pipe cross-sectional area to obtain the fluid dynamic viscosity; calculating the average flow velocity of the pipe in combination with the current rotational speed of the perfusion pump 9; multiplying the calculated average flow velocity of the pipe by the calculated fluid dynamic viscosity and dividing by the pipe characteristic scale factor to calculate the local shear stress inside the culture chamber 3; wherein, the pipe characteristic scale factor is an equivalent length parameter characterizing the hydraulic characteristics of the pipe cross-section, which can be obtained through calibration experiments with blank culture medium; In this invention, local shear stress is used to reflect the relative level of the fluid velocity gradient near the surface of the microcarrier. Its value cannot be directly obtained from a single flow signal. Therefore, it is calculated by combining the power-side signal and structural parameters. During pulse perfusion, the stepper motor 19 will undergo loading, stabilization and unloading processes. When the apparent viscosity of the culture medium increases, the time required for the motor torque fluctuation to recover to a steady state is prolonged. The controller 10 constructs a torque fluctuation curve by sampling the drive current, phase voltage, or out-of-step compensation amount, and defines the time taken from the torque peak to within 10% of the difference between the peak and steady-state value as the torque fluctuation decay time. The dynamic viscosity of the fluid is obtained by multiplying the decay time by the flow resistance coefficient related to the pipe cross-sectional area. The flow resistance coefficient is determined by the pipe inner diameter, effective length, peristaltic pump head 20 compression amount, and local resistance of the flow channel, and can be obtained by calibration with blank culture medium. Taking a system with an inner diameter of 2mm and an effective flow channel length of 500mm as an example, a flow resistance coefficient database can be established at a temperature of 37℃, and the controller 10 can call up the table; the rotation speed of the current irrigation pump 9 is determined by the pulse frequency of the stepper motor 19, and the discharge per revolution of the peristaltic pump head 20 is known, so the volumetric flow rate per unit time can be calculated, and then divided by the cross-sectional area of ​​the pipe to obtain the average flow velocity of the pipeline. Multiplying the average flow velocity by the calculated fluid dynamic viscosity and dividing by the pipeline characteristic scale factor yields the local shear stress used to characterize the local rheological load on the microcarrier surface; this conversion relationship is an engineering mapping in this invention, the purpose of which is to convert measurable quantities of electrodynamics into control quantities related to cell forces. Since an increase in extracellular matrix content in the culture medium leads to increased viscosity, the calculated results of local shear stress can reflect the shear risk during the high-density amplification phase and provide a basis for subsequently reducing the pulse flow rate. The above calculation process actually constitutes a local rheological and shear estimation model. The purpose of this model is to use the motor electrical signal on the power side to assess the shear risk exerted on the cells by the fluid near the surface of the microcarrier in real time under non-contact conditions. In terms of logical structure and data flow, the model includes a fluid dynamic viscosity calculation submodule and an average flow velocity estimation submodule; the former receives torque fluctuation decay time and called flow resistance coefficient as input and outputs fluid dynamic viscosity; the latter receives parameters such as stepper motor speed as input and outputs pipeline average flow velocity; the main model receives the output results of these two submodules and obtains local shear stress through division logic. In terms of the physical relationships represented, this model characterizes the damping effect of fluid viscosity dissipation on the motor load recovery process in a non-constant pulsed pipe flow system. Because the increased extracellular matrix secreted by cells in the culture medium leads to increased viscosity, which in turn prolongs the decay time of motor torque fluctuations, combining the torque fluctuation decay time with the average flow velocity can indirectly and effectively quantify the local shear risk on the surface of the microcarrier. The role of torque ripple decay time in the control logic is as an intermediate state quantity reflecting the fluid damping recovery rate. This intermediate state quantity is first used to estimate the fluid dynamic viscosity, and then the dynamic viscosity is used to participate in the calculation of local shear stress, rather than being directly used as shear strain rate. The preferred specific processing flow includes: Step 1, the controller 10 records the baseline current or baseline back electromotive force of the stepper motor 19 at the beginning of each pulse perfusion cycle; Step 2, continuously collects drive current or back electromotive force data during the high pulse flow rate input phase to obtain the torque peak moment; Step 3, continuously samples from the peak moment until the signal recovers to the steady state range to obtain the torque fluctuation decay time. Step 4: Controller 10 calls the pre-calibrated flow resistance coefficient table to map the decay time to the fluid dynamic viscosity; Step 5: Combine the current stepper motor 19 speed, peristaltic pump head 20 displacement per revolution and pipe cross-sectional area to obtain the average flow velocity; Step 6: Convert the average flow velocity and dynamic viscosity together into local shear stress and output it to the force field parameter scheduling module. The flow resistance coefficient is preferably determined through graded calibration, that is, the corresponding torque fluctuation decay time is measured under at least three sets of standard liquids with known viscosity, and a one-to-one correspondence between decay time and viscosity is established; when the system is changed to silicone hose 18 specification, peristaltic pump head 20 clamping amount or flow channel size, the corresponding flow resistance coefficient is recalibrated to ensure that the conversion diameter is consistent. The local shear stress obtained by multiplying the average flow velocity by the fluid dynamic viscosity and dividing by the pipeline characteristic scale factor is used as an equivalent index for control. Its physical meaning is to characterize the relative shear risk level exerted by the fluid near the microcarrier on the cell attachment interface under the current rheological conditions. An increase in the value of this index indicates that the flow driving ability under unit viscosity dissipation conditions is enhanced, and the probability of a high velocity gradient appearing on the surface of the microcarrier is increased simultaneously. To prevent individual pulses from causing calculation jumps, the controller 10 preferably takes a moving average of the calculation results over 3 to 10 consecutive pulse cycles and uses the smoothed result as the scheduling input. To further clarify the data flow and calculation rules of the above extrapolation process, a specific quantitative extrapolation example is provided below: Assume that the system uses a silicone tube 18 with an inner diameter of 2mm and a cross-sectional area of ​​approximately 3.14mm²; In a certain pulse perfusion, the controller 10 collects the drive current of the stepper motor 19 at a sampling rate of 10kHz. After identifying the torque peak, it records that the time taken for the current to decay to within 10% of the steady-state value is 45ms, which is the torque fluctuation decay time; The controller 10 calls the coefficient corresponding to the current pipeline configuration in the pre-stored flow resistance coefficient table, for example, 0.02 mPa·s / ms, and calculates the current fluid dynamic viscosity as 0.9 mPa·s through multiplication logic; at the same time, the controller 10 reads the current stepper motor 19 speed setting value, and combines it with the known peristaltic pump head 20 displacement per revolution, for example, 0.1 mL / rev, to calculate the current volumetric flow rate as 1.5 mL / s, and then divides it by the pipeline cross-sectional area to obtain the average pipeline flow velocity of approximately 477 mm / s; The controller 10 multiplies 477 mm / s by 0.9 mPa·s and divides it by the preset pipeline characteristic scale factor to calculate the equivalent characterization value of local shear stress as 530 mPa. This result is stored in the moving average cache queue of the controller 10 as the direct numerical basis for subsequent force field parameter scheduling. As can be seen from the above examples, this method does not require complex formulas. It can convert the underlying electrical signals of the motor into physical quantities that characterize the shear risk using only basic arithmetic logic, and has extremely high software programmability and real-time processing feasibility.

[0027] The inner wall of the inner liner 6 of the culture device is machined with an inverted conical blind hole 11, and the open end of the inverted conical blind hole 11 is sealed with a polyurethane elastic film 12 to form an inverted conical air chamber 13; the steps of dynamically scheduling the force field parameters of the transducer array 8 and the perfusion pump 9 based on the calculated cell confluence and local shear stress include: When the calculated cell confluence is less than the preset confluence threshold, the transducer array 8 is controlled to output the basic maintenance power, and the stepper motor 19 of the perfusion pump 9 is controlled to run at a high pulse width duty cycle and output a high pulse flow rate, so that the high pulse fluid squeezes the polyurethane elastic film 12 on the inner liner 6 into the inverted conical air chamber 13. In the early stages of cell expansion, the number of cells attached to the surface of the microcarrier is at the initial seeding level, and mass transfer limitation is greater than shear limitation. Therefore, the stage of enhanced perfusion is defined as when the cell confluence is less than a preset confluence threshold. The preset confluence threshold can be set from 20% to 60% depending on the cell type, preferably 40%. After the controller 10 identifies that the cell confluence is below the threshold, it controls the transducer array 8 to output the basic maintenance power. The basic maintenance power refers to the power level that is sufficient to maintain the existence of the standing wave structure in the culture chamber 3 without causing acoustic hindrance to the flow field that exceeds the predetermined rheological influence threshold. It is preferably 10% to 35% of the rated power. Meanwhile, the stepper motor 19 controlling the irrigation pump 9 operates with a high pulse width duty cycle, which can be 50% to 90%, so that the irrigation pump 9 maintains a high flow rate output time that meets the preset pulse width ratio in each pulse cycle, so as to promote the exchange of nutrients and gases through the microcarrier pores; the transient pressure generated by the high pulse flow rate acts on the polyurethane elastic film 12 on the inner liner 6, the film is concave into the inverted conical air chamber 13, the air chamber is compressed, the local volume increases, and the system absorbs part of the water hammer energy; This process is not an ancillary phenomenon, but rather part of the force field regulation. Without the aforementioned indentation energy absorption during the high pulse flow rate phase, the instantaneous pressure within the culture chamber 3 would fluctuate substantially, causing the node positions of the orthogonal ultrasonic standing wave field to shift and affecting the stability of the flow field near the microcarrier. By maintaining the combination of power and high pulse width duty cycle, as well as the synchronous response of the polyurethane elastic film 12 and the inverted conical air chamber 13, the mass transfer capacity within the pores can be improved in the early stages of amplification, while simultaneously controlling the pressure peak and local shear fluctuations.

[0028] A safety error envelope is set, which is a pre-defined local shear stress allowable range that meets the target cell viability requirements; the step of dynamically scheduling the force field parameters of transducer array 8 and perfusion pump 9 based on the calculated cell confluence and local shear stress also includes: When the calculated cell confluence is greater than or equal to the preset confluence threshold, the transducer array 8 is controlled to increase the transmission power, while the pulse width duty cycle of the perfusion pump 9 is shortened and the pulse flow rate is reduced. When the calculated local shear stress reaches the upper limit of the safety error envelope, the pulse width duty cycle is reduced or the transmit power of the transducer array 8 is increased to keep the local shear stress within the safety error envelope. When the cell confluence is greater than or equal to the preset confluence threshold, the culture system enters the high-density expansion stage, the cell layer thickens, the extracellular matrix increases, and the non-Newtonian characteristics of the culture medium are enhanced. If the initial high pulse flow rate is maintained, there is a risk that the local shear stress on the surface of the microcarrier may exceed the tolerance range of the cell membrane. To this end, the controller 10 increases the transmission power of the transducer array 8 to the enhanced locking power, preferably 35% to 80% of the rated power, to enhance the constraint effect of acoustic radiation on the cells and the surrounding fluid micro-regions. Based on the acoustic characteristics of standing wave fields, suspended cells and microcarriers are mainly constrained and concentrated in the nodal region with the smallest sound pressure amplitude under the action of acoustic radiation force, while the antinode region with the largest sound pressure amplitude becomes the channel with the least fluid resistance due to the lack of microcarrier obstruction. Therefore, after the acoustic radiation force is enhanced, a relatively stable high-viscosity acoustic retardation layer can be formed around the cell periphery, and the high-speed mainstream is more distributed in the antinode region, thereby reducing the shear load directly acting on the cell membrane surface. At the same time, the controller 10 shortens the pulse width duty cycle of the perfusion pump 9, which can be reduced to 10% to 50%, reducing the pulse flow rate and the duration of high flow rate; the controller 10 continuously compares the calculated local shear stress with the pre-established safety error envelope, which refers to the boundary of the allowable shear range calibrated based on cell type, microcarrier material and culture cycle. When the local shear stress approaches the upper limit of the envelope, the controller 10 continues to lower the pulse width duty cycle or increase the transducer power; when the local shear stress is lower than the median of the envelope and the cell confluence is still increasing, the controller 10 maintains the current parameters; by increasing the emission power and shortening the pulse width duty cycle of the perfusion pump 9 during the high-density amplification stage, the device can maintain the necessary mass transfer while keeping the shear stress on the cells within the safe error envelope, reducing abnormal shedding in the later stages of amplification; During this process, a force field safety scheduling model runs inside the controller 10. The purpose of this model is to balance mass transfer requirements with the risk of shear damage during the high-density cell expansion phase, and to prevent cells from falling off due to mechanical forces exceeding the damage threshold. In terms of logic structure and data flow, the model receives local shear stress and cell confluence as real-time inputs, and performs logical comparisons with the pre-stored safety error envelope and the preset confluence threshold. Finally, it outputs the transmit power adjustment command for the transducer array 8 and the pulse width duty cycle adjustment command for the perfusion pump 9. This model characterizes the protective mechanism of the locally high-viscosity acoustic retardation layer on cells induced by acoustic radiation force, as well as the weakening effect of pulse velocity reduction on the overall fluid shear load. Because high-power ultrasonic standing waves can confine the high-speed mainstream to the antinode region, and at the same time, reducing the pulse width duty cycle can directly weaken the fluid kinetic energy, the synergistic effect of the two can ensure that the local shear stress on the surface of the microcarrier is strictly limited within the safety error envelope. The preset confluence threshold in this invention serves as a trigger condition for the control strategy to switch from the enhanced perfusion stage to the low shear protection stage. Its determination is not arbitrary, but is set based on the inflection point of proliferation rate, the inflection point of shedding risk, and the change in metabolic consumption rate corresponding to the target cell type in the preliminary experiment. Preferably, multiple batches of culture experiments are first conducted under the same microcarrier loading and culture medium system, and the offline cell counting results, transducer resonant frequency drift, cell detachment ratio and nutrient consumption rate are recorded respectively; the interval where the risk of cell detachment begins to exceed the safe tolerance limit and further increasing the perfusion intensity no longer substantially improves the amplification efficiency is identified, and the lower limit of cell confluence corresponding to this interval is determined as the preset confluence threshold. For different cell types, the threshold can be set in the range of 20% to 60%, and the corresponding formula can be called by the controller 10 before the culture program starts; The physical meaning of the safety error envelope in the control logic is to characterize the allowable range of local shear stress that the target cell can tolerate during the current culture cycle. It includes at least three types of judgment boundaries: upper limit of the envelope, median of the envelope, and lower limit of the envelope. The upper limit of the envelope is used to trigger protective actions to prevent the cell membrane or attachment interface from being subjected to fluid shear exceeding the damage threshold; the median of the envelope is used to determine whether the current parameters can be maintained; the lower limit of the envelope is used to identify the risk of insufficient mass transfer and prevent nutrient exchange from falling below the metabolic requirement limit in order to protect the cell. The preferred steps for establishing the safety error envelope include: Step 1, selecting the target cell type, microcarrier material, and predetermined culture cycle as the calibration object; Step 2, running the culture device under different perfusion pulse width duty cycles and different transducer power combinations. Step 3: Record the corresponding local shear stress, cell viability, shedding rate, and amplification fold; Step 4: Define the range of local shear stress that meets the viability requirements and has a shedding rate lower than the preset upper limit as the safe range; Step 5: Further shrink the safe range according to the conservative principle to form a safe error envelope and store it in the controller 10 parameter library. During operation, the controller 10 preferably makes decisions in the following order: first, it compares the cell confluence with a preset confluence threshold. If the threshold has been reached or exceeded, it enters the high-density amplification stage. Then, it compares the positional relationship between the local shear stress and the safety error envelope. If the stress is higher than the upper limit of the envelope, it prioritizes shortening the pulse width duty cycle. If the stress is still higher than the upper limit of the envelope after shortening, it continues to increase the transducer power. If the value is near the median of the envelope, the current parameters are maintained; if the value is below the lower limit of the envelope and the cell confluence growth rate is below the preset growth threshold, the pulse width duty cycle is increased by a preset step size to restore mass transfer; where the preset growth threshold is the minimum confluence growth rate required to maintain basic cell proliferation as determined by previous experimental statistics; the preset step size is a fixed adjustment increment of the pulse width duty cycle. Therefore, the preset convergence threshold and the safety error envelope respectively serve the functions of stage switching judgment and fine adjustment judgment within the stage, and together constitute the logical basis for dynamic scheduling. To ensure that the above force field safety scheduling model has deterministic programmability at the software level, the specific rules for forming the safety error envelope by conservative shrinkage are further defined: After obtaining the safety interval, the controller 10 multiplies the upper limit of the interval by a scaling factor of 0.90 to 0.95 to obtain the upper limit of the envelope, multiplies the lower limit of the interval by a scaling factor of 1.05 to 1.10 to obtain the lower limit of the envelope, and takes the arithmetic mean of the upper and lower limits as the median of the envelope. The following is a quantitative deduction example of scheduling logic: Assume that the preset confluence threshold is set to 40%, the lower limit of the calibrated safety error envelope is 150mPa, the median is 300mPa, and the upper limit is 450mPa; During operation, the controller 10 receives that the current cell confluence is 45%. Since it is greater than 40%, the protection strategy of the high-density amplification stage is triggered. At this time, the controller 10 reads the current smoothed local shear stress as 480 mPa; since 480 is greater than the upper limit of the envelope 450, the scheduling model determines that there is a risk of shear damage and immediately outputs a command to reduce the pulse width duty cycle of the perfusion pump 9 from the current 60% to 40%; In the next evaluation cycle, if the local shear stress drops to 460 mPa, which is still higher than 450, the controller 10 will output a command to increase the transmit power of the transducer array 8 from 40% to 50% of the rated power to enhance the acoustic retardation layer protection. If the local shear stress drops to 320 mPa, which is close to the median of 300, the controller 10 outputs an instruction to maintain the current pulse width and power parameters. Through this explicit threshold comparison and step-by-step adjustment logic, the data interaction relationship between the modules of the system is clear and can be reliably executed by the industrial microprocessor through basic conditional branch statements, avoiding the unpredictability brought about by fuzzy control.

[0029] The embodiments of this application have been described above with reference to the accompanying drawings. However, this application is not limited to the specific embodiments described above. The specific embodiments described above are merely illustrative and not restrictive. Those skilled in the art can make many other forms under the guidance of this application without departing from the spirit and scope of the claims, and all of these forms are within the protection scope of this application.

Claims

1. A cell culture device for automatic cell proliferation, characterized in that, include: The frame (1) includes a lower partition (2); The culture chamber (3) is fixed in the center of the frame (1) by a flange (4). The culture chamber (3) includes a cylindrical quartz glass tube (5) and an inner liner (6), wherein the inner liner (6) is coaxially sleeved inside the cylindrical quartz glass tube (5). A transducer array (8) is fixed to the outer wall of the cylindrical quartz glass tube (5); a perfusion pump (9) is fixed on the lower partition (2) of the frame (1), the perfusion pump (9) includes a stepper motor (19), wherein the output end of the perfusion pump (9) is connected to the bottom of the culture chamber (3); The controller (10) is connected to control the operation of the transducer array (8) and the perfusion pump (9).

2. The cell culture device for automated cell expansion according to claim 1, characterized in that, The outer wall of the inner liner (6) is press-fitted with the inner wall of the cylindrical quartz glass tube (5). The inner wall of the inner liner (6) is arrayed with inverted conical blind holes (11), wherein the opening of the inverted conical blind holes (11) faces the center of the culture chamber (3).

3. The cell culture device for automatic cell expansion according to claim 2, characterized in that, The inner liner (6) is made of polyetheretherketone material. The cone angle of the inverted conical blind hole (11) is 50° to 70°. The opening end of the inverted conical blind hole (11) is hot-melt welded with a polyurethane elastic film (12). The thickness of the polyurethane elastic film (12) is 0.05 mm to 0.20 mm. The polyurethane elastic film (12) seals the interior of the inverted conical blind hole (11) to form an inverted conical air chamber (13) with initial air pressure.

4. The cell culture device for automatic cell expansion according to claim 1, characterized in that, The transducer array (8) includes piezoelectric ceramic transducers (14), which are fixedly connected to the outer wall of the cylindrical quartz glass tube (5) in orthogonal directions on the same horizontal plane by conductive silver paste (15). The piezoelectric ceramic transducers (14) form orthogonal ultrasonic emission pairs in pairs.

5. The cell culture device for automated cell expansion according to claim 1, characterized in that, The culture chamber (3) also includes a bottom end cap (16) and a top end cap (21). The bottom end cap (16) is machined with a liquid inlet channel (17). The liquid inlet channel (17) is connected to the output end of the perfusion pump (9) through a silicone hose (18). The perfusion pump (9) is directly driven by a stepper motor (19) to drive the peristaltic pump head (20). The top end cap (21) of the culture chamber (3) is machined with a liquid outlet channel (22).

6. A control method for a cell culture apparatus for automated cell proliferation as described in claim 1, characterized in that, include: S1. Control the perfusion pump (9) to continuously input pulsed fluid into the culture chamber (3), and control the transducer array (8) to establish an orthogonal ultrasonic standing wave field inside the culture chamber (3); S2. The current resonant frequency of the transducer array (8) and the torque fluctuation decay time of the stepper motor (19) are obtained through the controller (10). S3. Calculate the cell confluence based on the drift of the resonant frequency of the transducer array (8). S4. Calculate the local shear stress based on the torque fluctuation decay time of the stepper motor (19) of the irrigation pump (9); S5. Dynamically schedule the force field parameters of the transducer array (8) and the perfusion pump (9) based on the calculated cell confluence degree and the local shear stress.

7. The control method according to claim 6, characterized in that, The method for calculating cell confluence includes: obtaining the reference resonant frequency in the initial cell-free state; subtracting the currently measured resonant frequency from the reference resonant frequency in the initial cell-free state to obtain the frequency drift; and multiplying the frequency drift by a compensation coefficient related to the sound velocity of the medium to obtain the cell confluence.

8. The control method according to claim 6, characterized in that, The method for calculating local shear stress includes: multiplying the torque fluctuation decay time by the flow resistance coefficient related to the cross-sectional area of ​​the pipeline to obtain the fluid dynamic viscosity; and calculating the average flow velocity of the pipeline in combination with the rotational speed of the current irrigation pump (9). The calculated average flow velocity of the pipeline is multiplied by the calculated dynamic viscosity of the fluid and divided by the pipeline characteristic scale factor to calculate the local shear stress inside the culture chamber (3); wherein, the pipeline characteristic scale factor is an equivalent length parameter characterizing the hydraulic characteristics of the pipeline cross section, which can be obtained through a calibration experiment of blank culture medium.

9. The control method according to claim 6, characterized in that, The inner wall of the inner liner (6) of the culture device is machined with an inverted conical blind hole (11), and the opening end of the inverted conical blind hole (11) is sealed with a polyurethane elastic film (12) to form an inverted conical air chamber (13); the steps of dynamically scheduling the force field parameters of the transducer array (8) and the perfusion pump (9) according to the calculated cell confluence and the local shear stress include: When the calculated cell confluence is less than the preset confluence threshold, the transducer array (8) is controlled to output the basic maintenance power, and the stepper motor (19) of the perfusion pump (9) is controlled to run at a high pulse width duty cycle and output a high pulse flow rate, so that the high pulse fluid squeezes the polyurethane elastic film (12) on the inner liner (6) into the inverted conical air chamber (13).

10. The control method according to claim 9, characterized in that, A safety error envelope is set, which is a pre-set local shear stress allowable range that meets the target cell viability requirements; the step of dynamically scheduling the force field parameters of the transducer array (8) and the perfusion pump (9) based on the calculated cell confluence and the local shear stress further includes: When the calculated cell confluence is greater than or equal to the preset confluence threshold, the transducer array (8) is controlled to increase the transmission power, while the pulse width duty cycle of the perfusion pump (9) is shortened and the pulse flow rate is reduced. When the calculated local shear stress reaches the upper limit of the safety error envelope, the pulse width duty cycle is reduced or the transmission power of the transducer array (8) is increased so that the local shear stress remains within the safety error envelope.

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