A collagen bio-ink for one-step in-situ real-time 3D printing, preparation method and application

Abstract

The application belongs to the technical field of 3D printing biomaterials, and particularly relates to a collagen biomaterial ink for in-situ real-time 3D printing by one-step method as well as a preparation method and application thereof. The collagen biomaterial ink is prepared by mixing a collagen solution, methacrylic anhydride and a photoinitiator, and is cured under irradiation of blue light with a wavelength of 405 nm; the collagen biomaterial ink has good extrudability, and can be 3D printed to obtain a fine micron-level structure; the biological scaffold printed by the collagen biomaterial ink has good biocompatibility and bioactivity, and can significantly promote the repair of skin damage.

Description

TECHNICAL FIELD

[0001] The application belongs to the technical field of 3D printing collagen biomaterials, and particularly relates to a collagen biomaterial ink for in-situ real-time 3D printing by one-step method, a preparation method and application. BACKGROUND

[0002] Skin, as the first line of defense between the human body and the outside world, is easily damaged by the external environment. Skin damage can be caused by various reasons, including acute wounds such as mechanical damage, chemical damage, surgical wounds, and chronic wounds such as burns, infections, and diabetes. Full-thickness skin damage is prone to infection with various environmental bacteria, leading to prolonged inflammatory response and difficulty in wound healing. In order to improve the efficiency of wound healing, autologous transplantation, allogeneic transplantation and xenotransplantation methods have been studied, but there are risks of infection, damage to other soft tissues, immune rejection and other risks. Therefore, tissue-engineered skin is widely developed as a barrier for damaged skin, covering according to the degree of the wound and helping the wound to heal. Collagen accounts for 75% of the dry weight of human skin and is the main structural protein of the skin ECM. Due to its excellent biocompatibility, biodegradability and low immunogenicity, collagen is widely used in the preparation of tissue engineering regeneration scaffolds.

[0003] 3D bioprinting, as one of the most advanced manufacturing technologies, has the potential to manufacture complex and precise tissue engineering scaffolds. Compared with traditional tissue engineering manufacturing technologies, 3D bioprinting technology has the advantages of simplicity, efficiency, high precision, customization, etc. Extrusion bioprinting is one of the most commonly used 3D bioprinting technologies, which has been successfully applied in tissue regeneration engineering such as skin, bone, heart valve, cornea, etc. However, independent collagen solution is difficult to print, and currently strategies such as support bath, changing temperature and pH, adding cross-linking agent after extrusion, etc. are used for 3D printing.

[0004] The combination of collagen biomaterial ink and 3D bioprinting technology has great application potential in the field of tissue engineering. However, there are still many problems in the 3D printing of collagen: ① The method of introducing a support bath is complicated and has the problem of residual support bath; ② The method of rapidly changing temperature or pH has the risk of denaturation of collagen; ③ The addition of cross-linking agent in collagen may cause toxic side effects due to its residue; ④ The above methods cannot achieve in-situ real-time 3D bioprinting of collagen. Therefore, how to obtain a method for rapidly in-situ real-time 3D printing of collagen scaffolds is a technical problem that needs to be solved by those skilled in the art.

[0005] In view of the above technical problems, the present application surprisingly finds that after adding a trace amount of methacrylic anhydride (MA) to collagen, mixing with a photoinitiator LAP can realize efficient in-situ real-time 3D bioprinting of collagen by "one-step method". The collagen biological scaffold obtained by the above-mentioned 3D printing method has micron-level resolution, good mechanical strength, and anti-swelling and anti-degradation properties. The collagen biological scaffold has excellent biocompatibility and can promote the proliferation, adhesion, migration and differentiation of HFF-1 cells. In a rat full-thickness skin injury model, the collagen biological scaffold significantly promotes the repair of full-thickness skin injury. The development of the collagen biological ink "one-step in-situ real-time 3D bioprinting method has great application potential in the fields of tissue regeneration engineering and clinical medicine. SUMMARY

[0006] The primary object of the present application is to provide a "one-step in-situ real-time 3D bioprinting collagen biological ink, characterized in that the collagen biological ink is prepared by mixing collagen solution, methacrylic anhydride and photoinitiator, and is cured under the irradiation of blue light at a wavelength of 405 nm.

[0007] Preferably, the volume ratio of the collagen solution to the methacrylic anhydride in the collagen biological ink is 1:0.001-0.04.

[0008] Preferably, the volume ratio of the collagen solution to the photoinitiator solution is 1:0.01-1.

[0009] Preferably, the concentration of the collagen solution is 1-3% m / v.

[0010] Preferably, the photoinitiator is one or more of phenyl (2,4,6-trimethylbenzoyl) lithium phosphate (LAP), 2-hydroxy-1-[4-(2-hydroxyethoxy)phenyl]-2-methyl-1-propanone (I2959) and flavin mononucleotide (FMN).

[0011] Preferably, the photoinitiator is phenyl (2,4,6-trimethylbenzoyl) lithium phosphate (LAP).

[0012] The second object of the present application is to provide a method for 3D printing using the collagen biological ink, which comprises the following steps:

[0013] (1) preparing by dissolving freeze-dried collagen in an acetic acid solution;

[0014] (2) performing gradient dialysis on the collagen solution obtained in step (1) in an acetic acid solution;

[0015] (3) dialyzing the collagen solution after dialysis in step (2) in ultrapure water to neutral.

[0016] (4) adding methyl methacrylate and a photoinitiator into the collagen solution prepared in step (3) to prepare a collagen bio-ink;

[0017] (5) adding the bio-ink prepared in step (4) into a needle tube and placing the needle tube in an extrusion type 3D printer to perform in-situ real-time 3D printing by means of light curing after extrusion at 4-25 DEG C, so as to obtain a collagen bio-scaffold with a target structure.

[0018] A third object of the present application is to provide a collagen bio-scaffold prepared by the method.

[0019] A fourth object of the present application is to provide use of the collagen bio-scaffold in preparation of an implant, artificial skin, artificial bone, hemostatic sponge, scaffold material and medical device.

[0020] The present application has the following advantages: (1) the present application provides a collagen bio-ink for in-situ real-time 3D printing by one-step method, wherein the bio-ink is prepared by directly mixing methyl methacrylate, a photoinitiator LAP and a collagen solution;

[0021] (2) the collagen bio-ink has good and stable extrudability, and can be directly prepared into a target three-dimensional structure by extrusion type in-situ real-time 3D printing, thereby avoiding the risk of complicated steps and reduced activity of collagen caused by secondary cross-linking and maintaining stable shape after extrusion;

[0022] (3) the scaffold prepared from the collagen bio-ink has mechanical strength much higher than that of collagen with the same concentration, and has good mechanical strength, stable anti-swelling and anti-degradation performance;

[0023] (4) the scaffold prepared from the collagen bio-ink can significantly promote cell proliferation and adhesion;

[0024] (5) the scaffold prepared from the collagen bio-ink has excellent bioactivity, can significantly promote epidermal regeneration and ordered deposition of collagen fibers of full-thickness skin injury, and effectively promotes repair of skin injury. BRIEF DESCRIPTION OF DRAWINGS

[0025] Figure 1: Physicochemical properties of collagen (Col) and collagen bioink (CM-Bioink) Note: A is the apparent photo of the solidification of Col and CM-Bioink before and after light irradiation; B is the rheological results of Col and CM-Bioink before and after light irradiation; C is the apparent photo of the extrusion performance of CM-Bioink; D is the apparent photo of the uniformity of CM-Bioink; E is the push force of Col and CM-Bioink; F is the shear thinning behavior of Col and CM-Bioink; G is the gel-sol transition behavior of Col and CM-Bioink;

[0026] Figure 2 : Real-time extrusion 3D printing performance of CM-Bioink;

[0027] Figure 3 : Physicochemical properties of real-time extrusion 3D printed collagen scaffolds (CEPS) Note: A is the apparent photo of CEPS after freeze-drying; B is the SEM image of CEPS; C is the amplitude sweep results of Col and CEPS; D is the frequency sweep results of Col and CEPS; E is the compression curve of Col and CEPS; F is the compression modulus of Col and CEPS; G is the tensile curve of CEPS; H is the swelling behavior of CEPS; I is the degradation behavior of Col and CEPS;

[0028] Figure 4 : Cell compatibility and biological activity of CEPS Note: A is the cytotoxicity of CEPS to fibroblasts (HFF-1); B is the proliferation results of HFF-1 in CEPS; C is the live / dead cell staining image of HFF-1 cultured on CEPS for 1 day, 4 days and 7 days; D is the immunofluorescence staining image of HFF-1 cultured on CEPS for 1 day, 4 days and 7 days; E is the bright field image of CEPS promoting the migration of HFF-1; F is the migration rate of CEPS promoting the migration of HFF-1; G is the expression of myofibroblast-related genes alpha-smooth muscle actin (a-SMA), vimentin, collagen type I (Col-I) and collagen type III (Col-III) of HFF-1 cultured on CEPS for 7 days;

[0029] Figure 5 : Repair of full-thickness skin injury in rats by CEPS Note: A is the schematic diagram of in-situ real-time extrusion 3D printing of CEPS and its use for treatment of full-thickness skin injury in rats; B is the general observation photo of the wound healing process and the degree of skin repair; C is the wound healing rate;

[0030] Figure 6: CEPS histological staining results of rat full-thickness skin injury repair Note: A is H&E tissue staining figure; B is Masson tissue staining figure; C is epithelial regeneration rate; D is collagen deposition rate DETAILED DESCRIPTION

[0031] In order to make the technical means, creative features, purposes and effects realized by the present application easy to understand, the present application is further described below in conjunction with specific embodiments. However, the protection scope of the present application is not limited to the following embodiments.

[0032] The present application provides a "one-step" 3D-printed collagen biological scaffold with high precision, anti-swelling, anti-degradation, and high activity, which has excellent repair effect on full-thickness skin injury.

[0033] The collagen described in one or more of the following embodiments belongs to a biological polymer, is the main component of animal connective tissue, and is also the most abundant and widely distributed functional protein in the body of a mammal. It is formed by three polypeptide chains with left-handed helical structure winding around each other to form a right-handed helical structure.

[0034] The collagen described in one or more of the following embodiments can be natural collagen, recombinant collagen, and biomimetic collagen, as well as various naturally occurring or other ways of preparing collagen.

[0035] Example 1, preparation of collagen bio-ink (CM-Bioink)

[0036] 1.1 Preparation of collagen bio-ink

[0037] A 1M acetic acid solution was prepared, and 100mg of freeze-dried sponge-like collagen was dissolved in 10mL of acetic acid solution. After stirring and dissolving uniformly, it was placed in a 8000-14000 dialysis bag for dialysis. First, dialysis was performed in 0.5M acetic acid solution for 12h, then in 0.1M acetic acid solution for 12h, and finally in ultrapure water for 3 days. 10μL of MA was added to the collagen solution after dialysis to neutral, and 100μL of 10% LAP was added after mixing uniformly, to obtain collagen bio-ink (CM-Bioink) for extrusion 3D printing.

[0038] 1.2 Preparation of collagen bio-ink

[0039] A solution of acetic acid with a concentration of 1 M was prepared, 100 mg of freeze-dried sponge-like collagen was dissolved in 10 mL of the acetic acid solution, and after stirring and dissolving uniformly, it was placed in a dialysis bag with a molecular weight cut-off of 8000-14000 for dialysis. The dialysis was performed in a 0.5 M acetic acid solution for 12 h, in a 0.1 M acetic acid solution for 12 h, and finally in ultrapure water for 3 days. After the collagen solution was dialyzed to neutral, 40 μL of MA was added and mixed uniformly, and then 400 μL of 5% LAP was added to obtain the CM-Bioink for extrusion 3D printing.

[0040] 1.3 Preparation of collagen bio-ink

[0041] A solution of acetic acid with a concentration of 3 M was prepared, 300 mg of freeze-dried sponge-like collagen was dissolved in 10 mL of the acetic acid solution, and after stirring and dissolving uniformly, it was placed in a dialysis bag with a molecular weight cut-off of 8000-14000 for dialysis. The dialysis was performed in a 2 M acetic acid solution for 12 h, in a 1 M acetic acid solution for 12 h, in a 0.5 M acetic acid solution for 12 h, in a 0.1 M acetic acid solution for 12 h, and finally in ultrapure water for 3 days. After the collagen solution was dialyzed to neutral, 400 μL of MA was added and mixed uniformly, and then 10 mL of 0.2% LAP was added to obtain the CM-Bioink for extrusion 3D printing.

[0042] Example II, property study of collagen bio-ink (CM-Bioink)

[0043] 1. Collability of collagen bio-ink (CM-Bioink)

[0044] 10 mg / mL of collagen (Col) and the CM-Bioink described in Example 1.2 above were each placed in a test tube, and the gel state before and after irradiation at a 405 nm light source was observed by the test tube inversion method.

[0045] The results are shown in Figure 1 A. Before irradiation, both Col and CM-Bioink exhibited a milky white appearance. After 10 s of irradiation, Col did not change visibly and still had fluidity, while CM-Bioink solidified to form a stable hydrogel, indicating that CM-Bioink can be rapidly cross-linked under 405 nm irradiation.

[0046] 2. Rheological mechanics of collagen bio-ink (CM-Bioink) before and after irradiation

[0047] Multi-step rheological tests were used to quantify the dynamic mechanical properties of Col and CM-Bioink before and after irradiation. The changes in storage modulus (G') and loss modulus (G") of Col and CM-Bioink were measured at 25 °C before, 30 s after and after irradiation at 405 nm, keeping the angular frequency at 1 rad / s and the strain at 1%.

[0048] The results are shown in FIGS. Figure 1 As shown in FIG. B, CM-Bioink kept constant G' and G" before irradiation, and G' and G" increased significantly after 30 s irradiation, G' increased to 14 times before irradiation, while Col kept constant G' and G" before and after irradiation, indicating that the mechanical strength of CM-Bioink was significantly improved after irradiation at 405 nm.

[0049] 3. Extrudability of collagen bio-ink (CM-Bioink)

[0050] A needle with an inner diameter of 0.6 mm was selected as the standard to explore the extrusion test of CM-Bioink described in Example 1.2 above. CM-Bioink was placed in a 5 mL syringe, and the "Col" shape was extruded, accompanied by 405 nm blue light irradiation during the extrusion process, and a photo was taken.

[0051] The results are shown in FIGS. Figure 1 As shown in FIG. C, CM-Bioink presented a "Col" shape after extrusion, and solidification occurred immediately, indicating that CM-Bioink could be extruded into a specific shape, had good extrudability, and could maintain shape stability after extrusion.

[0052] To further verify the stability and uniformity of CM-Bioink, orange G was added to CM-Bioink and then extruded into the shapes of "butterfly" and "sun", accompanied by 405 nm blue light irradiation.

[0053] The results are shown in FIGS. Figure 1 As shown in FIG. D, CM-Bioink mixed with orange G still existed stably, and could continuously extrude uniform "butterfly" and "sun" shapes, indicating that CM-Bioink was a uniform sol that could be stably stored.

[0054] 4. Evaluation of the pushing force of collagen bio-ink (CM-Bioink)

[0055] To evaluate the size of the force required to extrude Col and CM-Bioink, they were placed in a 5 mL syringe used for 3D printing, the needle had an inner diameter of 0.25 mm, and the pushing force required by Col and CM-Bioink was detected using a needle tester at a room temperature and an extrusion rate of 10 mm / min, with time dependence (0-30 s).

[0056] Results are shown in FIG. 1A and FIG. 1B. Figure 1 As shown in FIG. 1C, the pushing force required for Col fluctuated significantly during the extrusion process, and the extrusion was unstable, while the CM-Bioink could be extruded stably as uniform filaments, and the pushing force required was about 15 N, indicating that the CM-Bioink had stable extrudability and was suitable for extrusion 3D printing.

[0057] 5. Shear-thinning behavior of collagen bio-ink (CM-Bioink)

[0058] The dynamic viscosity of Col and CM-Bioink was measured by a rheometer at 25°C in the shear rate range of 0.1-100 s -1 interval.

[0059] Results are shown in FIG. 2A and FIG. 2B. Figure 1 As shown in FIG. 2C, the viscosity of both Col and CM-Bioink decreased with the increase of shear rate, and the CM-Bioink decreased linearly, indicating that the CM-Bioink exhibited stable shear-thinning behavior in the test range and was suitable for extrusion 3D printing.

[0060] 6. Gel-sol transition behavior of collagen bio-ink (CM-Bioink)

[0061] Dynamic step strain amplitude tests were performed at 25°C under low strain (γ = 1.0%) and high strain (γ = 300%) conditions, and the extrusion process of Col and CM-Bioink in the syringe was simulated by continuous step changes in oscillatory strain.

[0062] Results are shown in FIG. 3A and FIG. 3B. Figure 1 As shown in FIG. 3C, after applying high strain (300%), the internal structure of the Col and CM-Bioink was immediately destroyed, resulting in a decrease in G' to about 9 Pa and 21 Pa, respectively, and G' was less than G", which was in a sol state. However, when converted to low strain (1%), the G' and G" of both were almost completely restored within a few seconds, and G' was greater than G", which was in a gel state. In three cycles of rupture and reorganization, both Col and CM-Bioink exhibited rapid sol-gel transition ability, indicating that they could rapidly convert to a sol state during the syringe extrusion process and immediately recover to the gel state before extrusion after extrusion.

[0063] Example Three, 3D printing of collagen bio-ink (CM-Bioink)

[0064] The CM-Bioink described in Example 1.2 was placed in an extrusion 3D printing needle tube, and the needle type was 25G (inner diameter 0.25 mm). A CAD model with a grid pattern of length 8.00 mm, width 8.00 mm, and height 2.00 mm was set and selected, with a line distance of 500 μm. The model was sliced, and the printing parameters were as follows: printing speed 3 mm / s, extrusion speed 0.3 mm / s, and the extruded profile was irradiated with a 405 nm light source immediately after the start of extrusion printing. After the end of printing a layer, the printed structure was observed and measured in size, and then a double-layer structure was printed to observe the effect of the cumulative number of printed layers on the printing size. 3

[0065] The single-layer printing result is shown in Figure 2 A, which shows that the CM-Bioink can be uniformly extruded and immediately cured after irradiation. Image J calculation shows that the line distance of the printed structure is (500.30 ± 10.42) μm, and the extruded line-shaped scaffold width is (262.63 ± 3.50) μm, which is consistent with the printing setting parameters; the double-layer printing result is shown in Figure 2 B, which shows that the double-layer printing does not cause structural collapse. The results show that the CM-Bioink can achieve extrusion real-time high-precision 3D printing, and the printed layers do not affect each other, and both single-layer printing and double-layer printing can be fine to the micron level structure.

[0066] Using the above parameters, a CAD model of a five-point star and a heart pattern was introduced, and the printing result is shown in Figure 2 C, which shows that high-efficiency 3D printing of three-dimensional structures can be successfully achieved; finally, a CAD three-dimensional model of a rabbit head and ears with complex structure was designed and introduced, and the printing result is shown in Figure 2 D, which shows that each printed structure accurately reflects the design in the CAD image, indicating that the CM-Bioink has successfully achieved 3D printing of complex three-dimensional structures.

[0067] Example Four, Physicochemical Properties of 3D Printed Collagen Porous Scaffold (CEPS)

[0068] 1. 3D Printed Collagen Porous Scaffold (CEPS)

[0069] The collagen bio-ink described in Example One was printed into a porous scaffold (CEPS) with a pore size of 1.00 mm under the 3D printing conditions described in Example Three. The printed CEPS was freeze-dried, as shown in Figure 3 A.

[0070] The micro-size of the above freeze-dried CEPS was characterized by scanning electron microscopy (SEM). After gold spraying treatment of the CEPS, it was scanned and measured at a voltage of 5.0 kV. ​

[0071] Results are shown in Figure 3 As shown in FIG. 2B, the SEM images presented a grid pattern with a grid length of (1.00 ± 0.05) mm, which was consistent with the grid model size, indicating that the bio-ink could achieve high-precision extrusion 3D printing.

[0072] 2. Mechanical properties of 3D printed collagen porous scaffolds (CEPS)

[0073] The mechanical properties of Col and CEPS were determined by rheological mechanics. The angular frequency was kept at 1 rad / s, and the amplitude sweep was performed in the strain range of 0.1%-100% to determine the linear viscoelastic region (LVR).

[0074] Results are shown in Figure 3 C, the CEPS had a wider LVR than the Col, and the G' of both LVRs was higher than the G". When the strain was 1%, the G' of the CEPS was 3974 Pa, while the G' of the Col was 467 Pa. The greater the G' value in the LVR region, the higher the hardness of the material. The results showed that both the Col and the CEPS behaved as viscoelastic solid in the LVR range, and the hardness of the CEPS was significantly higher than that of the Col.

[0075] When the shear strain was kept at 1%, the oscillatory shear rheological results with varying angular frequency are shown in Figure 3 D, the G' of the Col and the CEPS was always higher than the G" in the frequency range of 1-100 rad / s. The G' of the CEPS was almost unaffected by the frequency change, indicating that it remained in a stable gel state under short-time scale stress and long-time scale stress. The strength of the Col was significantly affected by the time scale, and the mechanical strength was significantly lower than that of the CEPS.

[0076] Compression tests were also used to evaluate the mechanical strength of the Col and the CEPS. The 3D printing prepared a cylindrical CEPS with a height of 2 mm and a diameter of 15 mm, and the Col was placed in the same size mold to prepare a cylinder. The hydrogel was placed on the lower compression plate, and the upper compression plate was set to move at a speed of 10 mm / min until it broke, and the force and displacement in each case were recorded.

[0077] The compression stress-strain curve is shown in Figure 3 E, the ultimate breaking strength of the CEPS was 216.68 kPa when the strain was 44.99%, and the compressive strength of the Col at this strain was 4.21 kPa. In the elastic stage (0-20%), the compression modulus results are shown in Figure 3 F, the compression modulus of the Col and the CEPS was (100.49 ± 3.46) kPa and (6.37 ± 1.88) kPa, respectively, with a significant difference. The above results showed that the mechanical strength of the CEPS was much higher than that of the Col at the same concentration.

[0078] Tensile measurements further confirmed the outstanding mechanical properties of CEPS. A 3D-printed rectangular CEPS was used, with a length, width, and thickness of 25 mm, 5 mm, and 1.5 mm, respectively. The tensile test speed was set to 10 mm / min. The elastic modulus was calculated using the slope of the stress-strain curve with an elongation of 25-30%.

[0079] The results are as follows Figure 3 As shown in Figure G, the maximum tensile stress of CEPS is 9.07 kPa, the strain is 33.72%, and its elastic modulus is (45.49 ± 2.50) kPa. These results all indicate that CEPS possesses excellent mechanical strength.

[0080] 3. Swelling behavior of 3D-printed collagen porous scaffolds (CEPS)

[0081] The swelling ratio (SR, %) of CEPS was determined by immersing it in deionized water at 25°C. The initial weight of the sample (W0, g) was obtained before immersion. After 1 hour, 2 hours, 4 hours, 8 hours, 18 hours, 24 hours, 30 hours, 39 hours, and 48 hours, the support was removed, and excess moisture on its surface was absorbed with filter paper. The sample was then weighed again (W1, g). The SR calculation formula is as follows:

[0082] SR(g / g)=W1 / W0×100%

[0083] The results are as follows Figure 3 As shown in H, the swelling of CEPS reached equilibrium within 8 hours, with a swelling ratio of approximately 113%. No visible changes occurred in the scaffold structure, indicating that CEPS has durable structural stability.

[0084] 4. Degradation behavior of 3D-printed collagen porous scaffolds (CEPS)

[0085] The degradation behavior of freeze-dried CEPS and the same mass of Col was evaluated in a 5 U / mL collagenase solution. The sample was accurately weighed (W0, mg), and 1 mL of collagenase solution was added and incubated at 37°C. The collagenase solution was changed daily to maintain enzyme activity. On days 1, 3, 5, 7, 9, 11, 13, and 15, the sample was removed from the enzyme solution, thoroughly washed three times with ultrapure water, centrifuged, and the supernatant was removed. The residual weight (W0) of the sample after freeze-drying was measured. t The degradation rate (MD, %) of the sample can be calculated using the following formula:

[0086] MD(%) = (W0 - W) t ) / W0×100%

[0087] The results are as followsFigure 3 As shown in Figure I, the degradation rate of Col was 81.06% on day 5, 95.60% on day 7, 97.20% on day 9, and completely degraded on day 11, while the degradation rate of CEPS was only 1.91% on day 15, indicating that CEPS has good anti-degradation performance in a short period of time.

[0088] Example 5: Cell compatibility and bioactivity of 3D-printed collagen bioscaffolds (CEPS)

[0089] 1. Cytotoxicity of 3D-printed collagen porous scaffolds (CEPS)

[0090] The cytotoxicity of CEPS prepared in Example 4 was assessed by detecting the viability of human foreskin fibroblasts (HFF-1) using the CCK-8 assay. HFF-1 cells were cultured in complete high-glucose medium (DMEM) containing 15% fetal bovine serum (FBS) and 1% penicillin-streptomycin antibiotics. 100 μL of the medium at a density of 1 × 10⁻⁶ μL was added to the corresponding wells of a 96-well plate. 5 mL -1 HFF-1 cell suspension was incubated in a cell culture incubator (37℃, 5% CO2) for 24 hours. After aspirating the supernatant, 100 μL of DMEM and CEPS extract were added to the corresponding wells. After incubation for 24 hours under the same conditions, 10 μL of CCK-8 was added and incubated for 1 hour. The absorbance at 450 nm was measured using a Tecan Infinite F200 / M200 multi-functional microwell analyzer.

[0091] The results are as follows Figure 4 As shown in Figure A, compared with the control group, the cell viability cultured in the CEPS extract was greater than 100%, indicating that CEPS has no cytotoxicity.

[0092] 2. Bioactivity of 3D-printed collagen porous scaffolds (CEPS)

[0093] The bioactivity of CEPS was evaluated using cell proliferation assays. The proliferation of HFF-1 cells in CEPS on days 1, 3, and 5 was assessed using the CCK-8 assay.

[0094] The results are as follows Figure 4 As shown in Figure B, on day 1 of culture, the relative proliferation rates of cells in the blank group and the CEPS group were 100.00% and 104.98%, respectively; by day 3, the proliferation rates increased to 104.12% and 111.78%, respectively; and by day 5, they reached 106.73% and 119.56%, respectively, indicating that CEPS can promote the proliferation of HFF-1 cells.

[0095] The cell survival status on CEPS hydrogel was further evaluated by live / dead cell staining. The CEPS was co-cultured with HFF-1 cell suspension, at each time point of day 1, day 4 and day 7, live cells and dead cells were stained with calcein-AM and propidium iodide respectively using live / dead cell staining kit at 37℃ for 1h, whether HFF-1 cells attached to CEPS and their survival status were observed.

[0096] The results of laser confocal scanning microscopy detection are shown in Figure 4 C, which shows that CEPS still maintains the original structure after 7 days of culture, and HFF-1 cells are uniformly distributed on the scaffold and grow throughout the scaffold at day 7. It shows that CEPS has excellent biological activity and can promote cell growth around and inside it.

[0097] The biological activity of CEPS was further explored by observing the adhesion of HFF-1 cells on CEPS. The CEPS was co-cultured with HFF-1 cell suspension in a laser confocal dish and incubated in a cell incubator (37℃, 5% CO2). The cell adhesion was observed after 1 day, 4 days and 7 days, and at each time point, the CEPS containing cells was sequentially fixed in 4% paraformaldehyde for 10 mins, permeated with 0.1% Triton X-100 for 5 mins, and then blocked with 1% BSA solution at room temperature for 30 mins. Subsequently, the actin cytoskeleton was stained with phalloidin-tetramethyl rhodamine at room temperature for 60 mins in the dark. Then, the cell nucleus was stained with Hoechst 33258 at 37℃ for 20 mins in the dark.

[0098] The fluorescence images taken by laser confocal microscope are shown in Figure 4 D, the HFF-1 cells adhered to CEPS are spindle-shaped and uniformly distributed, and show complex actin cytoskeleton structure, which shows that the biological scaffold provided by CEPS can significantly promote the adhesion and diffusion of HFF-1 cells.

[0099] The ability of CEPS to promote cell migration was evaluated by counting the change of cell scratch area over time. 2 mL of HFF-1 cells with a density of 1×10 5 mL -1 were placed in a six-well plate and incubated in an incubator (37℃, 5% CO2) for 24 hrs, then a blank area was artificially formed at the bottom of the well plate with a white gun head. After adding CEPS and incubating for 24 hrs, the cell migration in the well plate was observed by inverted fluorescence microscope and quantitatively analyzed by Image J software.

[0100] The results are shown in Figure 4As shown in Figure E, CEPS significantly promotes HFF-1 cell migration, with a migration rate of 80.29%. Figure 4 (as shown in F), while the blank group had only 26.86%.

[0101] After incubating CEPS cells in 6-well plates containing HFF-1 cells for 7 days, total RNA was extracted from the cells using an RNA isolation kit to analyze cell differentiation. First, the concentration and purity of the RNA samples were determined using a spectrophotometer. cDNA was synthesized using the PrimeScript RT kit and gDNA Eraser and analyzed using TB Green Premix Ex Taq II. Real-Time PCR was then performed using RT-qPCR to detect the expression of HFF-1 cell differentiation-related genes.

[0102] The results are as follows Figure 4 As shown in Figure G, compared with the control group, CEPS significantly promoted the expression of α-smooth muscle actin (α-SMA), vimentin, type I collagen (Col-I), and type III collagen (Col-III) in HFF-1 cells, indicating that CEPS can significantly promote the differentiation of fibroblasts into myofibroblasts. In conclusion, CEPS provides a highly bioactive scaffold for the adhesion, migration, and differentiation of HFF-1 cells.

[0103] Example 6: Repair of full-thickness skin damage in rats using 3D-printed collagen bioscaffolds (CEPS)

[0104] The repair effect of in situ real-time 3D printed collagen bioscaffolds (CEPS) on full-thickness skin injury was evaluated using a rat full-thickness skin injury model. Figure 5 (As shown in A). Sixty male SD rats (weighing 150–200 g) were purchased and anesthetized before modeling by intraperitoneal injection of 10% sodium pentobarbital (0.3 mL / 100 g). The rats' back hair was then shaved and disinfected with povidone-iodine. A circular area with a diameter of 10 mm was imprinted on the back, and the upper skin tissue was completely removed with surgical scissors, forming a circular full-thickness skin defect. The rats were randomly divided into a control group and an experimental group. The control group's wound was left untreated, while the experimental group underwent in-situ real-time 3D printing of CEPS (Continuous Extrusion Panel) that perfectly conformed to the wound model. The material was observed daily for detachment and wound healing. Samples were taken on days 4, 7, 14, 21, and 28 post-surgery, with 6 rats from each group euthanized each time. The wound edge tissue and granulation tissue were completely removed and fixed in 10% formaldehyde fixative for at least 3 days.

[0105] The wound healing was observed before each sampling, and the skin defect area was visualized and superimposed by Image J to clearly show the degree of skin repair. At the same time, the wound area was calculated by Image J, and then the wound healing rate was calculated.

[0106] The wound healing rate is shown in Figure 5 B, and the wound healing rate is shown in Figure 5 C. At day 4, the wounds of the control group and the CEPS group began to shrink, and the wound healing rates were 25.77% and 47.74%, respectively. At day 7, some scabs in the CEPS group naturally fell off, and the wound healing rate reached 85.12%, while the control group was 62.58%. At day 14, scabs fell off in both groups, and the wound healing rates of the control group and the CEPS group were 82.38% and 98.07%, respectively. At day 21, the wound closure rates of the control group and the CEPS group were 90.42% and 99.07%, respectively. At day 28, the wounds in the CEPS group were completely replaced by regenerated skin, while the control group still had 1.40% of residual wounds.

[0107] The fixed sample was gradient dehydrated in alcohol, embedded in paraffin, and sliced into 6 μm thin sections. Hematoxylin & eosin (H&E) was used for tissue section staining.

[0108] The staining results are shown in Figure 6 A, and the epithelial regeneration rate is shown in Figure 6 C. At day 4 after modeling, granulation tissue appeared in the CEPS group, and epithelial cells began to crawl. The epithelial regeneration rates of the control group and the CEPS group were 25.14% and 48.01%, respectively. At day 7, there were a large number of granulation tissues in the control group and the CEPS group, and the epithelial regeneration rates were 67.77% and 83.90%, respectively. By day 14, a thin epithelial layer formed on the surface of the CEPS group, and the epithelial regeneration rate reached 98.76%, while the control group was 84.35%. At day 21, the CEPS group began to grow reticular epithelium, and the maturity increased, with an epithelial regeneration rate of 99.45%. The control group delayed wound healing, and the epithelial regeneration rate was 91.64%. By day 28, the CEPS group achieved complete functional re-epithelialization, similar to normal skin. The middle part of the wound in the control group was not completely covered by the epithelium, and the epithelial regeneration rate was 97.79%. These experimental results verified that CEPS had a significant repair effect on full-thickness skin defects, and the newly formed regenerated skin showed similar dense and orderly characteristics to normal skin.

[0109] Masson's trichrome staining was used to show the deposition of collagen fibers in the skin, and the results are shown in Figure 6 B, and the collagen fiber deposition rate is shown in Figure 6As shown in D. On the 4th day, the control group and the CEPS group had sporadic blue collagen fibers on the skin wound, and the proportions were 1.15% and 7.39%, respectively; on the 7th day, the newly formed collagen fibers in the CEPS group increased significantly, and the proportion was 31.30%, while the collagen fibers in the control group were still not obvious, and the proportion was 4.79%; on the 14th day, the number of collagen fibers in the control group and the CEPS group was 30.84% and 68.90%, respectively; on the 21st day, the control group and the CEPS group both had a large number of newly formed collagen fibers, and the percentages were 49.84% and 85.71%, respectively; by the 28th day, the regenerated collagen fibers in the CEPS group were densely arranged, reaching 95.11%, while the collagen fibers in the control group were distributed more sparsely, and the proportion was 71.15%. The above results show that CEPS can promote the development of granulation tissue, accelerate reepithelialization, and guide the orderly deposition of collagen fibers, significantly promoting the repair of full-thickness skin defect wounds.

[0110] In summary, the present application provides a one-step in-situ real-time 3D printing collagen bio-ink, which is obtained by directly mixing methacrylic anhydride and photoinitiator LAP with collagen solution; has good and stable extrudable performance, can be directly prepared into a target three-dimensional structure by extrusion in-situ real-time 3D printing, avoids the complicated steps and the risk of reduced collagen activity caused by secondary cross-linking, and can maintain a stable shape after extrusion, printing fine micron-level structures, achieving high-precision and high-efficiency printing of collagen; the scaffold prepared from the collagen bio-ink has a mechanical strength much higher than that of collagen at the same concentration, has superior mechanical strength, stable anti-swelling and anti-degradation performance; can promote cell proliferation, significantly promote cell adhesion and diffusion; has good biocompatibility and outstanding biological activity, can significantly promote the epidermal regeneration of rat full-thickness skin damage and the orderly deposition of collagen fibers, and effectively promote the overall repair of skin damage.

Claims

1. A method for in situ real-time 3D printing of collagen scaffolds using collagen bio-ink, characterized in that, The collagen bio-ink is prepared by mixing a collagen solution, methacrylic anhydride and a photoinitiator, and is cured under irradiation of blue light at a wavelength of 405 nm. The method comprises the following steps: (1) dissolving lyophilized collagen in an acetic acid solution; (2) gradient dialysis of the collagen solution obtained in step (1) in an acetic acid solution; (3) dialysis of the collagen solution after step (2) in ultrapure water to neutral; (4) adding methacrylic anhydride and a photoinitiator to the collagen solution of step (3) and uniformly mixing to prepare a collagen bio-ink; (5) adding the bio-ink obtained in step (4) to a needle tube and placing it in an extrusion type 3D printer, and performing in-situ real-time 3D printing by light curing after extrusion at 4-25 ℃, to obtain a collagen bio-scaffold with a target structure.

2. The method of claim 1, wherein, The volume ratio of the collagen solution to the methacrylic anhydride in the collagen bio-ink is 1:0.001-0.

04.

3. The method of claim 1, wherein, The volume ratio of the collagen solution to the photoinitiator solution in the collagen bio-ink is 1:0.01-1.

4. The method of claim 1, wherein, The concentration of the collagen solution is 1-3% m / v.

5. The method of claim 1, wherein, The photoinitiator is one or more of phenyl (2,4,6-trimethylbenzoyl) lithium phosphate, 2-hydroxy-1-[4-(2-hydroxyethoxy)phenyl]-2-methyl-1-propanone and flavin mononucleotide.

6. The method of claim 5, wherein, The photoinitiator is phenyl (2,4,6-trimethylbenzoyl) lithium phosphate.

7. A collagen bio-scaffold prepared by the method of claim 1.

8. Use of the collagen bio-scaffold of claim 7 in the preparation of a medical device.

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