Self-deformable suspended myocardial tissue flexible force sensor and preparation method thereof

Abstract

The application discloses a self-deformation suspended myocardial tissue flexible force sensor and a preparation method thereof. The sensor comprises a rigid substrate, a flexible sensing element and a cell culture support. The flexible sensing element comprises a flexible substrate layer, a sensitive grid and a stress layer. The flexible substrate layer is connected with the rigid substrate. The stress layer is attached to the lower surface of the flexible substrate after being pre-stretched. The sensitive grid is arranged between the flexible substrate and the stress layer. A U-shaped cut is arranged along the outer periphery of the sensitive grid. The flexible sensing element at the U-shaped cut releases residual stress by the stress layer to form a curved cantilever. The cell culture support is arranged below the curved cantilever and is connected with the free end of the curved cantilever. The application converts the contraction deformation of myocardial tissue into the resistance change of a piezoresistive sensitive grid, so that the contraction force sensing of myocardial tissue is realized.

Description

TECHNICAL FIELD

[0001] The present application belongs to the field of biomedical technology, and particularly relates to a self-deforming flexible force sensor for suspended myocardial tissue and a preparation method thereof. BACKGROUND

[0002] To solve the ethical and cost problems of in-vitro models of heart diseases, myocardial chips (HoC) with similar morphology, gene expression and force- mechanical properties to natural tissues are expected to be able to construct highly mature engineered heart tissues. And, constructing cell scaffolds and sensors on myocardial chips to reproduce the in-vivo microenvironment of in-vitro myocardial tissue culture and monitor the dynamic behavior of myocardial tissue are of great significance to improve the maturity of in-vitro myocardial tissue. The contraction and relaxation deformation characteristics of in-vitro myocardial tissue are one of the important aspects of evaluating the physiological and pathological states of myocardial tissue. Therefore, developing sensors to evaluate the contraction function of in-vitro myocardial tissue can provide key basic data for cell / tissue physiology, myocardial toxicity testing, disease modeling and drug screening research.

[0003] At present, the contraction force sensing technology of myocardial tissue chips mainly relies on optical images and flexible structure deformation to monitor the mechanical signals of myocardial tissue, including video analysis, traction force microscopy (TFM), calcium imaging and other optical image technologies, which have the limitations of being difficult to quantify and long-term monitoring. Most of the existing technologies based on flexible structure deformation utilize the cell adhesion characteristics to construct sensitive elements on two-dimensional structures to monitor tissue contraction force, which has a large difference from the three-dimensional microenvironment in-vivo. The monitoring of the contraction force of suspended myocardial tissue is limited by the manufacturing process, and it is difficult to realize the three-dimensional structuring of the sensor. SUMMARY

[0004] The present application aims to overcome the shortcomings of the prior art and provides a self-deforming flexible force sensor for suspended myocardial tissue and a preparation method thereof, which solves the problem of three-dimensional structuring of the sensor in the background art.

[0005] One of the technical solutions adopted by the present application to solve its technical problems is to provide a self-deforming flexible force sensor for suspended myocardial tissue, which comprises a rigid substrate, a flexible sensing element and a cell culture scaffold. The flexible sensing element comprises a flexible base layer, a sensitive grid and a stress layer. The flexible base layer is connected to the rigid substrate. The stress layer is attached to the lower surface of the flexible base after being pre-stretched. The sensitive grid is arranged between the flexible base and the stress layer, and a U-shaped notch is arranged along the outer periphery of the sensitive grid. The flexible sensing element at the U-shaped notch forms a curved cantilever by utilizing the stress difference (release of residual stress of the stress layer). The cell culture scaffold is arranged below the curved cantilever and is connected to the free end of the curved cantilever.

[0006] In a preferred embodiment of the present application, a rigid support is further included for fixing and limiting the single end displacement of the cell culture scaffold.

[0007] In a preferred embodiment of the present application, the rigid substrate is provided with a hollow groove, the flexible substrate layer is attached to the lower surface of the rigid substrate, and the sensitive gate and the curved cantilever are located below the hollow groove. The rigid substrate and the flexible substrate layer are firmly connected by adhesion or bonding.

[0008] In a preferred embodiment of the present application, the surface of the rigid substrate is provided with a solder pad, which is in direct contact with the two ends of the sensitive gate and is electrically connected to the terminal of the two ends of the rigid substrate through the internal lead of the rigid substrate.

[0009] In a preferred embodiment of the present application, the surface of the flexible substrate layer is provided with a microchannel, and the sensitive gate is embedded in the microchannel.

[0010] In a preferred embodiment of the present application, the rigid substrate is provided with a port, the microchannel is connected to the port, and the port is used to perfuse the flexible substrate layer to form the sensitive gate.

[0011] In a preferred embodiment of the present application, the thickness of the flexible substrate layer and the stress layer is 0.02-0.5mm, but the thickness of the two is not consistent, so that in the deformation process of the curved cantilever caused by the deformation of the myocardial cells, the sensitive gate is subjected to compressive stress or tensile stress, thereby generating resistance change; the width of the U-shaped cut is 1-8mm, the length is 1-5mm, and the area of the cell culture scaffold is 3-25mm 2 .

[0012] In a preferred embodiment of the present application, the material of the flexible substrate layer and the stress layer is a stretchable flexible polymer of polydimethylsiloxane (PDMS), thermoplastic polyurethane (TPU) or silica gel, the material of the sensitive gate is graphite or carbon nanotube doped PDMS, conductive hydrogel or liquid metal, and the cell culture scaffold is an ordered fiber formed by polycaprolactone (PCL), thermoplastic polyurethane (TPU) or polylactic acid (PLA).

[0013] In a preferred embodiment of the present application, a data acquisition device is further included for real-time acquisition of the resistance of the two ends of the sensitive gate, and the data is exported to the processor for comparison with the sensor calibration data to obtain the myocardial tissue contraction force.

[0014] The second technical solution adopted by the present application to solve the technical problem is to provide a preparation method of a self-deformation suspended myocardial tissue flexible force sensor, comprising the following steps:

[0015] 1) Light-cured 3D printing or photolithography to prepare a mold, pouring liquid stretchable polymer, curing and demolding to prepare a flexible substrate layer and a stress layer;

[0016] 2) Surface adhesion of the flexible substrate layer to the rigid substrate surface; uniaxial stretching of the stress layer, adhesion of the stress layer and the flexible substrate layer; and formation of a sensitive grid between the stress layer and the flexible substrate layer;

[0017] 3) Cutting a U-shaped notch around the sensitive grid on the flexible sensing element, releasing residual stress in the stress layer to form a curved cantilever;

[0018] 4) Preparing a cell culture scaffold, and connecting the cell culture scaffold to the curved cantilever.

[0019] In a preferred embodiment of the present application, in step 1), the liquid stretchable polymer includes a stretchable flexible polymer of polydimethylsiloxane (PDMS), thermoplastic polyurethane (TPU) or silicone, and the step is: using light-cured 3D printing or photolithography to prepare a mold, pouring the liquid stretchable polymer into the mold, constant temperature curing at 60-120℃ for 30min-4h, demolding; further, forming a flexible substrate layer with microchannels on the surface;

[0020] In a preferred embodiment of the present application, in step 2), after the adhesion of the stress layer and the flexible substrate layer, the sensitive material is injected from the surface port of the rigid substrate into the microchannels of the flexible substrate layer in a pressure injection manner to form a sensitive grid; or, the sensitive grid is first formed on the surface of the flexible substrate layer using a direct writing process, and then the stress layer and the flexible substrate layer are adhered.

[0021] In a preferred embodiment of the present application, in step 2), the flexible substrate layer and the rigid substrate are firmly connected by using glue or plasma bonding, and the stress layer is stretched using a uniaxial stretching clamp, and the stress layer is firmly connected with the flexible substrate layer in the stretched state by plasma bonding.

[0022] In a preferred embodiment of the present application, in step 2), a syringe or a pressure pump is used to inject the sensitive material from the surface port of the rigid substrate into the microchannels of the flexible substrate layer to another port, and the port of the rigid substrate is sealed with sealant; the sensitive material includes PDMS doped with graphite or carbon nanotubes, conductive hydrogel, liquid metal.

[0023] In a preferred embodiment of the present application, in step 3), a U-shaped notch is cut around the outside of the sensitive grid using laser cutting or die cutting, and the opening of the U-shaped notch is directed parallel to the stretching direction of the stress layer in step 2), so that the flexible sensing element inside the U-shaped notch bends to form a three-dimensional curved cantilever, and the principle of residual stress difference between multiple film layers is used to realize the three-dimensionality of the two-dimensional flexible sensing element.

[0024] In a preferred embodiment of the present application, in step 4), a rigid support is fixed below the connecting end of the curved cantilever, and the rigid support is bonded by using glue or plasma bonding, so that one end of the cell culture scaffold is connected with the rigid support to limit single-end displacement, and the other end is connected with the free end of the curved cantilever.

[0025] In a preferred embodiment of the present application, in step 4), a pair of grounding electrodes are placed on both sides of the rigid support and the curved cantilever, and ordered fibers are prepared by using the opposite electrode electrospinning, and the excess fibers are removed by using the hot cutting method to form the cell culture scaffold.

[0026] Compared with the background art, the technical solution has the following advantages:

[0027] The present application realizes the three-dimensional structuring of the two-dimensional flexible sensing element by using the principle of the difference of residual stress between the multiple film layers, can convert the contraction deformation of the myocardial tissue into the resistance change of the piezoresistive sensing grid, and simultaneously considers the suspension culture of the in-vitro myocardial tissue. The present application has the advantages of good biocompatibility, reproduction of the in-vivo microenvironment of the in-vitro myocardial tissue culture, high sensitivity, adjustable mechanical properties, mechanical matching with the myocardial tissue, and the like, and the preparation process of the present application is simple. BRIEF DESCRIPTION OF DRAWINGS

[0028] Figure 1 It is a structure schematic diagram of the flexible force sensor of the self-deforming suspension myocardial tissue of Example 1.

[0029] Figure 2 It is an exploded schematic diagram of the flexible force sensor of the self-deforming suspension myocardial tissue of Example 1.

[0030] Figure 3 It is a schematic diagram of the measurement device of the flexible force sensor of the self-deforming suspension myocardial tissue of Example 1.

[0031] Figure 4 It is a flow schematic diagram of the preparation method of the flexible force sensor of the self-deforming suspension myocardial tissue of Example 2.

[0032] Among them, 1 is a rigid substrate, 2 is a flexible base layer, 3 is a sensing grid, 4 is a stress layer, 5 is a rigid support, and 6 is a cell culture scaffold. DETAILED DESCRIPTION

[0033] The technical solution of the present application will be described clearly and completely in combination with the drawings and examples. It should be noted that the terms "upper", "lower", "inner", and the like indicate the orientation or positional relationship shown in the drawings, and are only used for the convenience of describing the present application and simplifying the description, and do not indicate or imply that the devices or elements referred to must have a particular orientation, be constructed and operated in a particular orientation, and therefore cannot be understood as a limitation on the present application.

[0034] Example 1

[0035] The self-deformation flexible force sensor for suspended myocardial tissue in this embodiment, as shown in Figure 1 and Figure 2 , comprises a PCB rigid substrate 1, a flexible sensing element and a cell culture scaffold 6; the flexible sensing element is composed of a flexible substrate layer 2 and a stress layer 4, the stress layer is attached to the lower surface of the flexible substrate after pre-stretching, the surface of the flexible substrate layer is provided with micro-channels, and a graphite-doped PDMS sensitive grid 3 is embedded therein; a U-shaped notch is arranged along the outer periphery of the sensitive grid 3, the flexible sensing element releases residual stress through the U-shaped notch to form a curved cantilever; the concave side of the curved cantilever (the connecting end of the curved cantilever) is provided with a rigid support 5, and is located at the lower surface of the flexible sensing element; one end of the cell culture scaffold 7 is connected with the rigid support 5, and the other end is connected with the free end of the curved cantilever, and the main body of the cell culture scaffold 7 is suspended below the curved cantilever and the rigid support, which takes into account the suspended culture of the in-vitro myocardial tissue.

[0036] The flexible substrate layer 2 is firmly bonded with the PCB rigid substrate 1 by plasma in a state without additional deformation, the PCB rigid substrate 1 is provided with a rectangular hollow slot of 6mm×15mm in the middle, the stress layer 4 is firmly bonded on the surface of the flexible substrate layer 2 by plasma bonding process in a uniaxial stretching state, and the U-shaped notch is cut on the surface of the flexible sensing element by laser processing to realize the curved cantilever. The rectangular hollow slot is used to prevent the flexible substrate layer 2 from being bonded with the rigid substrate 1 in this area, to ensure the effective inspection of the flexible sensor element, and also can be used to observe cells through this area.

[0037] As shown in Figure 1 , the myocardial tissue is inoculated, adhered and grown on the cell culture scaffold 7, when the myocardial tissue contracts, the cell culture scaffold deforms, the end face of the free end of the curved cantilever is displaced, and the sensitive grid 3 in the curved cantilever is strained, which causes the resistance to change.

[0038] In this embodiment, the thicknesses of the flexible substrate layer 2 and the stress layer 4 are 0.3mm and 0.05mm respectively, the thicknesses of the two are inconsistent, which ensures the existence of the stress difference required for self-deformation, the width of the U-shaped notch is 4mm, the length is 4mm, and the width of the cell culture scaffold 7 is 4mm, and the length is 6mm.

[0039] As shown in Figure 4 , 50μL of suspension containing about 0.5 million cells is injected above the cell culture scaffold 7 by using a pipette, and is placed in a culture dish, after 6 hours of cell sedimentation on the scaffold surface, culture solution is added until the cells are completely immersed. The sensor is placed in a cell incubator for culture until the cells expand and adhere to the cell scaffold.

[0040] As shown in Figure 3As shown, the surface of the rigid PCB substrate 1 has pads that directly contact the two ends of the sensitive gate 3, and are electrically connected to the terminals at both ends of the rigid PCB substrate 1 through internal PCB wires. The data acquisition device collects the resistance at both ends of the sensitive gate in real time, exports the data to a computer, and compares it with the sensor calibration data to determine the magnitude of the myocardial tissue contractile force.

[0041] Example 2

[0042] like Figure 4 As shown, this embodiment also provides a method for preparing the self-deformable suspended myocardial tissue flexible force sensor of Example 1, including the following steps:

[0043] 1) After uniformly mixing PDMS monomer and curing agent at a mass ratio of 10:1, the mixture is poured into a 3D-printed male mold with a sensitive grid structure. The mold is then placed in a vacuum chamber to remove any residual air bubbles in the PDMS. The mold is then placed in a constant temperature chamber and cured at 120°C for 30 minutes. Demolding yields a flexible substrate layer 2 with microchannels.

[0044] 2) Coat the surface of the rigid PCB substrate 1 with a layer of PDMS prepolymer and cure it in a constant temperature oven at 120℃ for 30 minutes. Place the flexible substrate 2 and the rigid PCB substrate 1 into a plasma cleaner and perform oxygen plasma cleaning at 300W for 80 seconds. After removal, bond the flexible substrate 2 to the surface of the rigid PCB substrate 1. The ports of the microchannels on the surface of the flexible substrate 2 are aligned with the ports on the surface of the rigid PCB substrate 1.

[0045] 3) Cut a 0.05mm thick PDMS film, slightly larger than the flexible substrate 2, as the stress layer 4. Use a uniaxial stretching fixture to stretch the film and place it in a plasma cleaner at the same time as the flexible substrate 2 bonded to the surface of the PCB rigid substrate 1. Then perform bonding according to step 2).

[0046] 4) Using an injection molding compound, 20% by mass of graphite-doped PDMS prepolymer was injected into the microchannel through one of the ports on the surface of the rigid PCB substrate 1 until it overflowed from the other port. The device was then placed in an oven and cured at 120°C for 30 minutes to form the sensitive gate 3.

[0047] 5) Use a femtosecond laser to cut a 4mm×4mm U-shaped cut along the periphery of the sensitive grid. The U-shaped cut is oriented parallel to the tensile direction of the stress layer 4. The flexible base layer 2, sensitive grid 3 and stress layer inside the U-shaped cut bend upward due to residual stress, forming a cantilever.

[0048] 6) Use a mold to prepare a 2mm high PDMS cube as a rigid support 5, and use PDMS prepolymer as glue to adhere it to the stress layer 4 on the concave side of the curved cantilever.

[0049] 7) Aluminium counter electrodes are placed on both sides of the rigid post 5 and the curved cantilever, the counter electrodes are connected to a negative voltage of -2 kV. A 20% mass fraction PCL solution in glacial acetic acid is placed in a syringe, the syringe is connected to a positive voltage of 8 kV and electrospinning is performed at a rate of 400 μl / h on the rigid post 5 and the curved cantilever. The electrospun fibres are deposited in an ordered fashion on the rigid post 5 and the curved cantilever under the influence of the counter electrodes to form an ordered fibre, excess fibre is removed using a hot knife to form a cell culture scaffold 7.

[0050] The above examples are only used to illustrate the technical solutions of the present application, not limit it; although the present application has been described in detail with reference to the foregoing examples, those of ordinary skill in the art should understand that they can still modify the technical solutions recorded in the foregoing examples, or make equivalent replacement for part or all of the technical features; and these modifications or replacements do not make the essence of the corresponding technical solutions deviate from the scope of the technical solutions of the embodiments of the present application.

Claims

1. A self-deformable flexible force sensor for suspended myocardial tissue, characterized in that: The system includes a rigid substrate, a flexible sensing element, and a cell culture scaffold. The flexible sensing element comprises a flexible substrate layer, a sensitive grid, and a stress layer. The flexible substrate layer is connected to the rigid substrate. The stress layer is attached to the lower surface of the flexible substrate after pre-stretching. The sensitive grid is disposed between the flexible substrate and the stress layer, and a U-shaped cut is provided along the outer periphery of the sensitive grid. The flexible sensing element at the U-shaped cut utilizes stress differences to form a bending cantilever. The cell culture scaffold is disposed below the bending cantilever and connected to the free end of the bending cantilever. It also includes rigid struts for fixing and limiting one-end displacement of the cell culture scaffold; The rigid substrate is provided with a hollowed-out groove, the flexible substrate layer is attached to the lower surface of the rigid substrate, and the sensitive gate and the bending cantilever are located below the hollowed-out groove; The surface of the flexible substrate layer is provided with microchannels, and the sensitive gate is embedded in the microchannels; The flexible substrate and stress layer are made of polydimethylsiloxane, thermoplastic polyurethane, or stretchable flexible polymers of silicone. The sensitive gate is made of graphite or carbon nanotube-doped PDMS, conductive hydrogel, or liquid metal. The cell culture scaffold is made of ordered fibers formed from polycaprolactone, thermoplastic polyurethane, or polylactic acid.

2. The self-deformable suspended flexible force sensor for myocardial tissue according to claim 1, characterized in that: The thickness of the flexible base layer and the stress layer is 0.02–0.5 mm, but their thicknesses are not the same; the width of the U-shaped incision is 1–8 mm, the length is 1–5 mm, and the area of ​​the cell culture scaffold is 3–25 mm². 2 .

3. The method for preparing a self-deformable, suspended flexible force sensor for myocardial tissue as described in claim 1 or 2, characterized in that: Includes the following steps: 1) Use photopolymerization 3D printing or photolithography to prepare molds, pour liquid-phase stretchable polymers, and cure and demold to prepare flexible substrate and stress layers; 2) Attach the surface of the flexible substrate layer to the surface of the rigid substrate; A uniaxial tensile stress layer is used to bond the stress layer and the flexible substrate layer; and a sensitive grid is formed between the stress layer and the flexible substrate layer. 3) Cut a U-shaped cut around the sensitive grid on the flexible sensing element and use the stress layer to release residual stress to form a bending cantilever; 4) Prepare a cell culture scaffold, placing it below and connecting it to the curved cantilever. In step 3), a U-shaped cut is cut around the outside of the sensitive grid using laser cutting or die cutting. The opening of the U-shaped cut faces the direction of the stress layer tension in step 2), so that the flexible sensing element inside the U-shaped cut bends to form a three-dimensional bending cantilever. In step 4), a pair of grounded electrodes are placed on both sides of the rigid support and the curved cantilever, and ordered fibers are prepared by electrospinning of the electrodes. Excess fibers are removed by thermal cutting to form a cell culture scaffold.

4. The method for preparing a self-deformable suspended flexible force sensor for myocardial tissue according to claim 3, characterized in that: In step 2), after the stress layer and the flexible substrate are bonded together, the sensitive material is injected from the surface port of the rigid substrate into the microchannel of the flexible substrate by pressure injection to form a sensitive gate; or, a direct writing process is first used to form a sensitive gate on the surface of the flexible substrate, and then the stress layer and the flexible substrate are bonded together.

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