Low-temperature DLP 3D printing method for low-concentration collagen and application of low-temperature DLP 3D printing method

Abstract

The invention belongs to the technical field of 3D printing, and particularly relates to a low-temperature DLP 3D printing method for low-concentration collagen and application of the low-temperature DLP 3D printing method. The printing performance of the photosensitive collagen ColMA bio-ink can be obviously improved in a pH 3 system for the first time, and the durable high-activity collagen biological scaffold is obtained through low-temperature DLP 3D printing; compared with a high-concentration GelMA scaffold, the low-concentration ColMA biological scaffold has higher mechanical performance, shorter equilibrium swelling time, lower swelling ratio, higher degradation resistance and higher biological activity, can remarkably promote proliferation, adhesion and migration of human prepuce fibroblasts and differentiation of the human prepuce fibroblasts into myofibroblasts, and has the advantages of being capable of remarkably improving the human prepuce fibroblasts in the human prepuce fibroblasts and improving the human prepuce fibroblasts in the human prepuce fibroblasts in the human prepuce fibroblasts in the human prepuce fibroblasts in the human prepuce fibroblasts. The inflammatory response of the rat full-thickness skin injury part is effectively relieved, wound reepithelization and ordered collagen deposition are promoted, full-thickness skin injury regeneration is accelerated, and wide application prospects are achieved in tissue regeneration engineering.

Description

Technical Field

[0001] The present invention belongs to the field of collagen DLP 3D bioprinting, and particularly relates to a low-temperature DLP 3D printing method for low-concentration collagen and its application. Technical Background

[0002] 3D bioprinting is a new type of biofabrication technology aimed at reconstructing human tissues and organs. Traditional tissue engineering methods cannot achieve precise bionic manufacturing and replication of complex cell microenvironments, which greatly limits their application in the field of regenerative medicine. 3D bioprinting can achieve personalized bionics of tissue structures, and has important advantages such as fast printing speed, short preparation time, and high resolution, and is expected to solve challenges in biological research and clinical applications. Currently, various forms of 3D bioprinting technologies have been developed, and the most common ones are inkjet bioprinting and extrusion bioprinting. Inkjet bioprinting has relatively low costs and medium printing speeds, but has the disadvantage of nozzle clogging when using high-viscosity materials and high-density cells. Extrusion printing is suitable for bioinks of various viscosities. However, the printing time is long, and it is difficult to achieve high-precision bioprinting. Digital light processing (DLP) bioprinting is a light-assisted 3D printing technology, which has advantages such as fast printing speed, high printing accuracy, and no nozzle clogging problems, overcomes the defects of inkjet bioprinting and extrusion bioprinting, and has attracted much attention in biological application fields such as disease modeling and regenerative medicine.

[0003] The DLP 3D printing bioink needs to simultaneously possess stable fluidity, high photopolymerization rate, sufficient mechanical strength, and excellent biocompatibility to achieve precise replication of natural tissues. Currently, the types of bioinks used for DLP 3D printing are limited, and they are mainly prepared by introducing a photoreactive part into natural biomaterials, such as silk fibroin, hyaluronic acid, and gelatin. Methacrylated gelatin (GelMA) has advantages such as low cost, simple preparation, and good biocompatibility, and is one of the most widely used DLP 3D printing bioinks in the field of tissue regeneration. However, as a denatured collagen, gelatin lacks the triple-helix structure of collagen, and has defects such as fast degradation rate, poor mechanical properties, and low biological activity, which greatly limits its application in tissue engineering and regenerative medicine. Therefore, it is urgent to develop new high-activity DLP 3D printing bioinks.

[0004] Collagen is an indispensable component of the extracellular matrix, and its unique triple-helix structure makes it not degraded by most proteases. In recent years, collagen-based materials have been widely used in biomedicine and tissue engineering. Collagen has advantages such as good biocompatibility, low immunogenicity, appropriate biodegradability, and excellent biological activity, and has great application potential as a 3D printing bioink.

[0005] Collagen is the main component of many organs in the human body. The bioink based on collagen has important application prospects in the tissue and organ fabrication by DLP 3D bioprinting. The ε-amino group on lysine of collagen can react with methacrylic anhydride to synthesize photoreactive collagen ColMA through amidation reaction, enabling photocrosslinking between collagens to form hydrogels. Currently, there are still the following problems regarding DLP 3D printing of collagen: ① If the 3D printing time is too long, the bioink outside the sliced area of the model will solidify, resulting in the edge of the printed structure being blurred and the printing resolution being reduced; ② Most collagen bioinks have a risk of denaturation during the preparation process, leading to a decrease in the bioactivity of the collagen scaffolds fabricated by 3D printing; ③ When using undenatured collagen as the bioink, the mechanical strength of the medical materials obtained by printing is poor, and it is impossible to accurately print well-defined and structurally stable medical tissues / organs; ④ Currently, collagen is usually combined with synthetic polymers such as PEGDA for DLP 3D bioprinting, but its limited biocompatibility and the potential toxicity of its degradation by-products hinder its wide application in the biomedical field. Summary of the Invention

[0006] In view of the above technical problems, the present invention unexpectedly discovers that by dissolving ColMA in an acetic acid solution to a pH of 2.5 - 3.0 and mixing it with the photoinitiator LAP, a highly active collagen scaffold can be prepared by DLP 3D printing mechanism at low temperature. Compared with the high-concentration GelMA scaffold printed by DLP 3D at high temperature, the low-concentration ColMA scaffold has stronger mechanical properties, as well as significantly enhanced anti-swelling performance and durability. Moreover, the low-concentration ColMA scaffold has more excellent bioactivity compared to the high-concentration GelMA scaffold and can significantly promote the repair of full-thickness skin wounds in rats. Based on the development of low-concentration collagen bioink for DLP 3D bioprinting at low temperature, it has broad application prospects in the fields of tissue engineering and biomedicine. Specifically, it includes the following contents:

[0007] In the first aspect, the present invention provides a method for low-temperature DLP 3D printing of ColMA biomaterials with low-concentration collagen, and the method is as follows: Dissolve the photosensitive collagen ColMA in an acetic acid solution, then mix it evenly with the photoinitiator LAP solution, and perform DLP 3D printing at 4 - 25 °C under blue light irradiation with a wavelength of 405 nm to obtain a collagen biomaterial; the ColMA is collagen modified with methacrylic anhydride.

[0008] Preferably, the method includes the following steps:

[0009] (1) Obtain the photosensitive collagen ColMA by methacrylation modification of collagen;

[0010] (2) Dissolve the photosensitive collagen ColMA obtained in step (1) in an acetic acid solution to obtain a photosensitive collagen ColMA solution;

[0011] (3) Add a photoinitiator LAP to the photosensitive collagen ColMA solution described in step (2) to prepare a collagen bioink;

[0012] (4) Add the bioink obtained in step (3) to a DLP 3D printer with a blue light wavelength of 405 nm, and perform layer-by-layer printing by photocuring molding at 4 - 25 °C. The curing time for each layer is 1 - 5 s;

[0013] (5) Subject the material formed by printing in step (4) to secondary curing at 4 - 25 °C under 405 nm blue light, and the curing time is 1 - 8 s, then the ColMA biomaterial printed by DLP 3D can be obtained.

[0014] Preferably, the pH of the ColMA solution is 2.5 - 3.0.

[0015] Preferably, the pH of the ColMA solution is 3.0.

[0016] Preferably, the concentration ratio of the ColMA solution to the LAP solution is 1:0.1 - 5.

[0017] Preferably, the concentration of the ColMA solution is 0.1 - 1% m / v.

[0018] Preferably, the final concentration of the LAP solution is 0.1 - 0.5% m / v.

[0019] Preferably, the preparation method of the ColMA is as follows: dissolve collagen in an acetic acid solution, adjust the pH of the reaction system to 7 - 9, then add an excessive amount of methacrylic anhydride, react for 24 hrs under ice bath conditions, dialyze and purify, and obtain the photosensitive collagen ColMA after freeze-drying.

[0020] Preferably, the photosensitive collagen ColMA is stored at -20 °C for standby.

[0021] Preferably, the preparation method of the ColMA includes the following steps:

[0022] (1) Dissolve freeze-dried collagen in an acetic acid solution to prepare a collagen solution;

[0023] (2) Adjust the pH of the collagen solution described in step (1) to 7 - 8 with 0.1 M sodium hydroxide solution, add an excessive amount of methacrylic anhydride dropwise, and obtain a reaction product after reacting for 24 hrs under ice bath conditions;

[0024] (3) The reaction product obtained in step (2) is dialyzed with ultrapure water for 4-8 times to obtain a ColMA solution.

[0025] Preferably, the concentration of the collagen stock solution is 0.01-1% m / v.

[0026] Preferably, the concentration of the collagen stock solution in the step is 0.1-0.5% m / v.

[0027] Preferably, the molar ratio of the collagen to methacrylic anhydride is 1:60-100.

[0028] Preferably, the method for extracting collagen comprises the following steps:

[0029] Pretreatment: Wash and crush the yak tendon tissue to obtain small pieces of tissue material;

[0030] Decellularization: small pieces of tissue were sequentially passed through NaCl solution, chlorhexidine solution and H 2 O 2 The solution removes cells and some residual impurities;

[0031] Degreasing: Degrease the decellularized tissue with n-butanol solution 2-4 times;

[0032] Decalcification: Decalcify the defatted tissue with hydrochloric acid solution 1-2 times;

[0033] Removal of foreign proteins and endotoxins: Remove the hydrochloric acid solution and soak the decalcified tissue in 0.01-2.0M sodium hydroxide solution for 2-24 hours;

[0034] Extraction of collagen stock solution: remove the sodium hydroxide solution, extract collagen with acetic acid-pepsin solution, purify the extract by salting out, and dialyze with acetic acid solution and ultrapure water for 6-8 times in sequence to obtain collagen stock solution.

[0035] In a second aspect, the present invention provides a ColMA biomaterial prepared by printing using the method described in the first aspect.

[0036] In a third aspect, the present invention provides the use of the ColMA biomaterial described in the second aspect in the preparation of a skin damage repair product.

[0037] The beneficial effects of the present invention are as follows: ① By directly dissolving collagen ColMA modified with methacrylic anhydride in an acetic acid system and adding a photoinitiator LAP, a collagen bioink that can be used for low-temperature DLP 3D printing is obtained, which is simple to operate and easy to prepare; ② The collagen bioink provided by the present invention maintains the complete triple helix structure of collagen. The bioink has good fluidity and can be ultra-rapidly cured and formed within 5 s under blue light irradiation, meeting the characteristic requirements of DLP 3D printing; ③ The present invention provides a low-temperature DLP 3D printing method using collagen alone as a bioink, realizing the broad-spectrum printing of collagen; ④ The present invention also provides a low-concentration collagen scaffold for DLP 3D printing. This low-concentration ColMA scaffold has higher mechanical strength, a smaller swelling ratio, and significant anti-degradation performance compared to a high-concentration GelMA scaffold; ⑤ The low-concentration ColMA scaffold prepared by low-temperature DLP 3D printing provided by the present invention has higher biological activity than a high-concentration GelMA scaffold and shows a more excellent repair effect in a rat full-thickness skin injury model. Brief Description of the Drawings

[0038] Figure 1 The fluidity, curability of the bioink ColMA of the present invention, and the DLP 3D printing grid scaffold and its printing accuracy;

[0039] Figure 2 The fluidity, curability, and DLP 3D printability of the bioink of ColMA of the present invention at different pH values;

[0040] Figure 3 The precise contours of various models printed by low-temperature DLP 3D printing of the bioink ColMA of the present invention;

[0041] Figure 4 Comparison of the physicochemical properties between the low-temperature DLP 3D printed low-concentration ColMA scaffold and the high-temperature DLP 3D printed high-concentration GelMA scaffold of the present invention;

[0042] Figure 5 Comparison of the biological activities between the low-temperature DLP 3D printed low-concentration ColMA scaffold and the high-temperature DLP 3D printed high-concentration GelMA scaffold of the present invention;

[0043] Figure 6 Comparison of the rat full-thickness skin injury repair effects between the low-temperature DLP 3D printed low-concentration ColMA scaffold and the high-temperature DLP 3D printed high-concentration GelMA scaffold of the present invention;

[0044] Figure 7Histological comparison of the repair of full-thickness skin defects in rats using low-concentration ColMA scaffolds printed by low-temperature DLP 3D printing and high-concentration GelMA scaffolds printed by high-temperature DLP 3D printing in the present invention. Detailed implementation mode

[0045] To make the technical means, creative features, achieved purposes and functions of the present invention easy to understand, the present invention will be further described below in conjunction with specific implementation modes. However, the protection scope of the present invention is not limited to the following examples.

[0046] The low-concentration ColMA scaffold printed by low-temperature DLP 3D printing provided by the present invention has higher mechanical strength and more excellent biological activity than the high-concentration GelMA scaffold printed by high-temperature DLP 3D printing, and can significantly promote the regeneration of full-thickness skin defects in rats.

[0047] In the embodiments of the present invention, only yak collagen is used as the raw material for the preparation of the 3D printing bioink, but the present invention is not limited to yak collagen, and any collagen obtained by other means can be used, including type I collagen, type II collagen, type III collagen, etc.

[0048] Example 1 Preparation of bioink ColMA

[0049] Dissolve the freeze-dried collagen in an acetic acid solution (0.3% m / v), stir to dissolve evenly, then adjust the pH to 7 with 0.1M NaOH solution, add 10 times the excess methacrylic anhydride, and keep the system pH at 7-8. React for 24 hrs under ice bath conditions, and dialyze and purify with ultrapure water to obtain collagen ColMA with photoreactive groups. After freeze-drying, store it at -20°C for later use.

[0050] Before DLP 3D printing, dissolve ColMA in an acetic acid solution to a final concentration of 0.5% (m / v), adjust the system pH to 3, and add the photoinitiator LAP to a final concentration of 0.2% (m / v), and mix evenly to obtain the ColMA bioink.

[0051] Example 2 DLP 3D printing of bioink ColMA

[0052] Place the ColMA bioink described in Example 1 in a vial, and observe its fluidity by the method of inverting the vial. The experimental phenomenon shows that the bioink presents a uniform and clear appearance and has good fluidity ( Figure 1 as shown in A). After irradiating with a 405nm light source for 3s, the bioink immediately solidifies and loses its fluidity, and a stable hydrogel is formed at the bottom of the vial ( Figure 1as shown in B). The changes in the mechanical strength of ColMA bioink before and after photopolymerization were evaluated by rheological tests. The frequency-dependent oscillatory shear rheology results showed that the storage modulus (G') of ColMA hydrogel was always higher than the loss modulus (G") in the frequency range of 1-100 rad / s, about 505 Pa, while the G' of ColMA bioink was about 55 Pa ( Figure 1 as shown in D). Strain-dependent oscillatory rheology indicated that the critical strain value of ColMA hydrogel at the gel-sol transition point was about 79%, while the critical strain value of ColMA bioink was about 37% ( Figure 1 as shown in E). The micro-morphology of the above gels was characterized by scanning electron microscopy (SEM) ( Figure 1 as shown in C). After sputtering gold on the freeze-dried ColMA gel, it was scanned and measured at a voltage of 5.0 kV. The SEM image presented a three-dimensional interconnected porous network structure with an average pore size of (95.66±15.55) μm. This porous structure could promote cell attachment, proliferation, nutrient transport, and metabolic waste excretion.

[0053] To demonstrate the low-temperature DLP printing performance of ColMA bioink and study the optimal printing parameters, a CAD model of a grid pattern with a length of 13.00 mm, a width of 13.00 mm, and a height of 3.00 mm was designed, and the side length of the internal square pores was 2.00 mm ( Figure 1 as shown in G). The above collagen bioink was introduced into a DLP 3D printer, and then the above model was selected and sliced. The printing parameters were as follows: the light intensity was 5 mW / cm 2 , the number of base layers was 5, the exposure time of the base layer was 4 s, the exposure time of each layer was 3.8 s, the temperature was 16 °C, and DLP 3D printing was carried out under 405 nm blue light irradiation; after printing, a blue light flashlight with a wavelength of 405 nm was used as a light source and placed 1 cm away from the printing support, and irradiated at 16 °C for 5 s. After secondary curing, a ColMA bioink DLP 3D printed grid scaffold was obtained ( Figure 1 as shown in H). An inverted microscope was used to observe the microscale size of the DLP 3D printed ColMA hydrogel grid ( Figure 1 as shown in F), and the image showed that the side length of the square mesh holes in the ColMA hydrogel was (2.01±0.07) mm. SEM was used to characterize the microstructure of the freeze-dried DLP 3D printed ColMA hydrogel grid ( Figure 1As shown in I). The SEM image presents a square grid pattern with a grid side length of (2.00 ± 0.11) mm. All the measured dimensions are consistent with the grid model dimensions, indicating that the DLP 3D printing of ColMA bioink can achieve precise manufacturing of three-dimensional structures, has excellent printability, and can be used as a separate bioink for DLP 3D printing.

[0054] Example 3 Fluidity, Curing Property, and DLP 3D Printability of ColMA Bioink under Different pH Conditions

[0055] Disperse the ColMA described in Example 1 in different pH solutions, and study the fluidity, curing property, and DLP 3D printing performance of 0.5% ColMA bioink in systems with pH 2, pH 3, pH 4, pH 5, pH 6, and pH 7. Take 3 mL of each and place them in different vials. Observe the fluidity of each bioink by the vial inversion method. The experimental phenomena show that all bioinks have good fluidity, but only the bioinks in the pH 2 and pH 3 systems present a uniform and clear appearance. After adding 0.3% LAP and irradiating with a 405 nm light source for 3 s, only in the pH 3 system does the bioink immediately solidify to form a stable hydrogel. Under the printing conditions described in Example 2, only the bioink in the pH 3 system can print a clear heart shape ( Figure 2 as shown).

[0056] Example 4 Broad Spectrum of DLP 3D Printing of Bioink ColMA in pH 3 System

[0057] In addition, verify the broad spectrum of ColMA as a DLP 3D printing bioink in the pH 3 system by DLP 3D printing three-dimensional structures of various colors and shapes under the printing parameters described in Example 2. We printed various complex shapes, such as a heart ( Figure 3 as shown in A), a plum blossom ( Figure 3 as shown in B), a leaf ( Figure 3 as shown in C), and an ear ( Figure 3 as shown in D). Each printed structure accurately reflects the design in the CAD image, demonstrating the applicability of ColMA bioink for DLP 3D printing in the pH 3 system. To make more diverse geometric patterns using ColMA bioink in DLP 3D printing, we uniformly mixed various dyes, including riboflavin, coomassie brilliant blue, methylene blue, rhodamine B, phenol red, orange G, brilliant green SF, and hematoxylin. The above bioinks precisely prepared uniform scaffolds in the following shapes, including a five-pointed star ( Figure 3 as shown in E), a peach shape ( Figure 3 as shown in F), the outline of a fish ( Figure 3 as shown in G), a petal shapeFigure 3 as shown in H), rabbit outline ( Figure 3 as shown in I), butterfly outline ( Figure 3 as shown in J), honeycomb structure ( Figure 3 as shown in K) and LZU model ( Figure 3 as shown in L). Further, it is shown that ColMA can be used as a stable bioink for DLP 3D printing to replicate various three-dimensional structures under a pH 3 system. The broad-spectrum nature of ColMA as a DLP 3D printing bioink provides broad application prospects in the field of tissue engineering.

[0058] Example 5 Comparison of the mechanical strength between low-temperature DLP 3D printed low-concentration ColMA scaffolds and high-temperature DLP 3D printed high-concentration GelMA scaffolds

[0059] Gelatin is a denatured product of collagen and does not have the triple-helix structure of collagen. Therefore, a gelatin derivative GelMA with photoreactive groups was used as a control group to study the related properties of ColMA. Using the printing parameters described in Example 2, a 0.5% ColMA hydrogel scaffold (ColMA-S) was DLP 3D printed at 16 °C and a 5% GelMA hydrogel scaffold (GelMA-S) was DLP 3D printed at 37 °C.

[0060] At 25 °C, an Antonpa rheometer was used to measure the mechanical properties of ColMA-S and GelMA-S. When the shear strain was kept constant at 1%, a scan was performed in the frequency range of 0.1 - 100 Hz to determine the frequency-dependent rheological properties ( Figure 4 as shown in A). Within the frequency range, the value of G' was always greater than G", and the G' values of ColMA-S and GelMA-S were approximately 4318 Pa and 928 Pa, respectively. This result indicates that both scaffolds exhibited elastic gel behavior and showed solid-like characteristics consistent with the viscoelastic behavior of typical gels. Notably, ColMA-S exhibited significantly superior mechanical properties. The viscoelastic properties of ColMA-S and GelMA-S were further revealed by strain-dependent oscillatory rheology with a constant angular frequency of 1 rad / s ( Figure 4 as shown in B). Within the linear viscoelastic region (LVR), the G' of both scaffolds was not affected by the applied strain and showed characteristics similar to those of viscoelastic solids. Using the G' value within the LVR as an index, the higher the G' value, the greater the stiffness of the scaffold. Within the LVE range, the G' value of ColMA-S was 4.67 times that of GelMA-S.

[0061] At 25 °C, the dynamic viscosities of ColMA-S and GelMA-S were measured in the shear rate range of 0.01 - 100 s -1 interval ( Figure 4As shown in C. The results show that as the shear rate increases, the viscosity decreases, indicating that both scaffolds exhibit shear thinning behavior within the test range. The strength of intermolecular interactions was evaluated by calculating the slope of the viscosity fitting curve. A larger slope indicates stronger interactions. The slope of the viscosity curve of ColMA-S was 0.8948, and the slope of the viscosity curve of GelMA-S was 0.8784. The results show that both scaffolds exhibit properties similar to viscoelastic gels, similar to the viscoelastic properties inherent in the extracellular matrix (ECM) and most biological tissues. This result emphasizes the potential use of the above 3D printed scaffolds in highly biomimetic tissue structures.

[0062] The compression test is a test to determine the mechanical properties of materials under axial static pressure and is one of the basic methods for material mechanical property tests. The mechanical properties of ColMA-S and GelMA-S were further evaluated through compression tests. Uniaxial compression tests were performed on ColMA-S and GelMA-S hydrogel cylinders (diameter 13 mm, height 2.5 mm) using a universal testing machine with a load cell to characterize the mechanical properties. The hydrogel was placed on the lower platen, and the upper platen was set to move at a speed of 10 mm / min until rupture, and the force and displacement in each case were recorded. Comparing the compression stress-strain curves of ColMA-S and GelMA-S, when the strain was 59.10%, the ultimate fracture strength of ColMA-S was 118.98 kPa, while when the strain was 56.08%, the ultimate fracture strength of GelMA-S was 14.07 kPa ( Figure 4 As shown in D. In the elastic stage (0 - 20%), their compression moduli were (12.51 ± 1.17) kPa and (2.49 ± 1.25) kPa respectively ( Figure 4 As shown in E. The preloading force for the cyclic compression test was set to 0.005 N, and then the test was started. The loading speed was 10 mm / min. The specimen was compressed to 40% strain and then returned to the initial position for a total of 1 cycle ( Figure 4 As shown in F. Hysteretic energy is a direct indicator of the internal energy dissipation of the hydrogel. Compared with GelMA-S, ColMA-S exhibited a larger hysteretic area, confirming its higher energy dissipation efficiency. These results further highlight the significant advantage of ColMA-S over GelMA-S in terms of compressive resistance.

[0063] Example 6 Comparison of the Swelling Behaviors of Low-Concentration ColMA Scaffolds Printed by Low-Temperature DLP 3D Printing and High-Concentration GelMA Scaffolds Printed by High-Temperature DLP 3D Printing

[0064] The ColMA-S and GelMA-S prepared in Example 5 were immersed in deionized water to measure their swelling ratio (SR) ( Figure 4As shown in Figure G. First, carefully weigh the initial weight of the sample (W 0 , g), and then immerse it in deionized water at 25°C. Every 2 hours, 4 hours, 8 hours, 12 hours, 24 hours, 36 hours, and 48 hours, take the sample out of the water, carefully remove the excess water on its surface with filter paper, and weigh the sample (W 1 , g). The calculation formula for the swelling ratio (SR, %) is as follows:

[0065] SR (g / g) = W 1 / W 0 × 100%

[0066] Compare the swelling properties of the two hydrogel scaffolds. The results show that the swelling ratio of ColMA-S is around 98%, and its swelling behavior reaches equilibrium within 2 hrs. While the swelling ratio of GelMA-S increases significantly, and the equilibrium swelling ratio is as high as 800%, and the hydrogel shows a visibly multiplied swelling state. ColMA-S shows a shorter equilibrium swelling time and a lower swelling ratio, indicating that ColMA-S has excellent dimensional stability.

[0067] Example 7 Comparison of the Degradation Degrees of Low-Concentration ColMA Scaffolds Printed by Low-Temperature DLP 3D Printing and High-Concentration GelMA Scaffolds Printed by High-Temperature DLP 3D Printing

[0068] We simultaneously evaluated the in vitro degradation properties of ColMA-S and GelMA-S prepared in Example 5 using trypsin and pepsin. Freeze-dry the ColMA-S and GelMA-S hydrogel scaffold samples, and accurately weigh the freeze-dried samples (W 0 , mg). Treat them separately with PBS solutions containing 5 U / mL pepsin and trypsin in different wells of a 24-well plate. Incubate at 37°C, and change the pepsin and trypsin solutions every day to maintain their enzyme activities. On the 1st day, 2nd day, 3rd day, 5th day, and 7th day, take the samples out of the enzyme solutions, thoroughly rinse them three times with ultrapure water, then centrifuge and remove the supernatant. Measure the remaining weight of the samples (W t , mg) after freeze-drying. The degradation rate (MD, %) of the samples can be calculated by the following formula:

[0069] MD (%) = (W 0 - W t ) / W 0 × 100%

[0070] In the trypsin solution, the degradation rate of GelMA-S is 100% on the 5th day, and the degradation rate of ColMA-S is less than 20%, and it still shows good stability on the 7th day ( Figure 4as shown in Figure H). In the pepsin solution, GelMA-S was basically completely degraded on the second day, and the degradation rate of ColMA-S was less than 20%, and it still showed good stability on the seventh day ( Figure 4 as shown in Figure I). The fast degradation rate of gelatin is an inevitable shortcoming, which may limit its application in tissue repair in vivo. The low-concentration ColMA DLP3D printed scaffold has good anti-degradation performance.

[0071] Example 8 Comparison of the bioactivity of low-temperature DLP 3D printed low-concentration ColMA scaffolds and high-temperature DLP 3D printed high-concentration GelMA scaffolds

[0072] 8.1 Cell compatibility of ColMA scaffolds and GelMA scaffolds

[0073] The cell compatibility of the ColMA-S and GelMA-S extracts prepared in Example 5 was evaluated by a cytotoxicity experiment. The in vitro cytotoxicity of the ColMA-S extract was determined by detecting the viability of human foreskin fibroblasts (HFF-1) using CCK-8, with the GelMA-S extract as the control group ( Figure 5 as shown in Figure A). HFF-1 cells were cultured in high-glucose medium (DMEM) supplemented with 15% fetal bovine serum (FBS) and 1% penicillin-streptomycin antibiotics, and cultured at 37°C in a 5% CO 2 atmosphere. Add 100 μL of HFF-1 cell suspension with a density of 1×10 5 mL -1 to a 96-well plate, and then place it in an incubator for 24 hrs. After aspirating the supernatant, add 100 μL of DMEM, ColMA-S, and GelMA-S extracts to the corresponding wells respectively. After incubating for 24 hrs, add 10 μL of CCK-8 to each well and incubate in the incubator for 1 hr, and then measure the absorbance at a wavelength of 450 nm using a Tecan Infinite F200 / M200 multi-functional microplate reader. Compared with the blank group, the cell viability in the ColMA-S extract and the GelMA-S extract was greater than 100%, indicating that neither of them had cytotoxicity. 8.2 Bioactivity of ColMA scaffolds and GelMA scaffolds

[0074] The cell activity of ColMA-S and GelMA-S was evaluated by a cell proliferation experiment. The CCK-8 method was used to evaluate the viability of HFF-1 cells in ColMA-S and GelMA-S hydrogels on the 1st, 3rd, and 5th days, and their proliferation was closely monitored ( Figure 5On the first day of culture, the relative proliferation rates of the blank group, GelMA-S group, and ColMA-S group were 100.00%, 100.31%, and 101.38%, respectively; on the third day, the proliferation rates increased to 120.62%, 123.82%, and 130.48%, respectively, and the difference between the blank group and the ColMA-S group was statistically significant; on the fifth day, the proliferation rates further increased to 130.30%, 141.45%, and 151.24%, respectively. Statistical analysis (p value) confirmed that ColMA-S could promote the proliferation of HFF-1 cells, which was significantly higher than the cell proliferation rate in GelMA-S.

[0075] The cell proliferation of ColMA-S and GelMA-S was further evaluated by live / dead cell staining ( Figure 5 (as shown in C). HFF-1 cells were attached to ColMA-S and GelMA-S for culture, and live / dead cell staining was performed on the 1st, 4th, and 7th days. Confocal laser scanning microscopy analysis showed that ColMA-S remained intact after 7 days of culture, and HFF-1 cells were evenly distributed on the scaffold. On the 7th day, the cells filled the entire scaffold. The ColMA-S printed scaffold has good cell compatibility, making it easy for cells to adhere and grow around and inside it. However, GelMA-S degraded during the culture process, resulting in the inability of cells to adhere and grow in sheets.

[0076] We used HFF-1 cells to perform cell adhesion experiments to evaluate the biological activities of ColMA-S and GelMA-S ( Figure 5As shown in D). HFF-1 cells were attached to ColMA-S and cultured, incubated in a laser confocal dish at 37°C. After 1 day, 4 days, and 7 days, 4% formaldehyde cell fixative was added to fix the cells for 10 mins, and the cells were permeabilized with 0.1% Triton X-100 for 5 mins. Then, they were blocked with 1% BSA in PBS buffer (10 mM, pH 7.2 - 7.4) at room temperature for 30 mins. The BSA solution was aspirated, and 100 nM phalloidin solution was directly added, and the cells were incubated in the dark at room temperature for 60 mins to stain the actin cytoskeleton. 5 μg / ml Hoechst 33258 solution was added, and the cells were incubated in the dark at 37°C for 20 mins to stain the cell nuclei, and then stored by soaking in PBS solution. Fluorescence observation was performed under a laser confocal microscope. The cell nuclei were observed using 360 nm excitation (blue fluorescence), and the cytoplasm was observed using 540 nm excitation (red fluorescence). HFF-1 cells adhered to ColMA-S showed a complex actin cytoskeleton structure. These cells were evenly distributed and showed a spindle-shaped morphology. In contrast, the cell distribution on GelMA-S appeared sparse. The experimental results clearly showed that compared with GelMA-S, the biocompatible scaffold provided by low-concentration ColMA-S could significantly enhance the adhesion and spreading of HFF-1 cells.

[0077] We evaluated the ability of ColMA-S and GelMA-S extracts to promote cell migration by statistically analyzing the change rate of the cell scratch area over time ( Figure 5 As shown in E). 2 mL of HFF-1 cells with a density of 1×10 5 mL -1 were placed in a six-well plate for culture. After culturing in an incubator at 37°C and 5% CO 2 for 24 hrs, a blank area, called a "scratch / wound", was artificially formed on the monolayer cells at the bottom of the well plate using a white pipette tip. When ColMA-S and GelMA-S hydrogel extracts were added, the HFF-1 cells at the edge received the migration signal and gradually entered the blank area to heal the "scratch / wound". After 24 hrs of migration, the migration of cells in each group of materials was observed through an inverted fluorescence microscope and quantitatively analyzed using Image J software. After 24 hrs of culture, the ColMA-S extract significantly promoted the migration of HFF-1 cells, and the migration rate was higher than 80% ( Figure 5 As shown in F), almost twice that of the GelMA-S extract.

[0078] We used HFF-1 cells for cell differentiation experiments. ColMA-S and GelMA-S were placed in 6-well plates adhered to HFF-1 cells and incubated for 1 day and 7 days. Total RNA was extracted from the above cultured cells using an RNA isolation kit. The concentration and purity of the RNA samples were determined using a spectrophotometer. cDNA was synthesized using the PrimeScript RT reagent kit and gDNA Eraser, and the expression level of the gene was analyzed by using TB Green Premix Ex Taq II. We used the Real-Time PCR system for RT-qPCR experiments to detect the expression of genes related to HFF-1 cell differentiation. Compared with GelMA-S, ColMA-S can significantly promote the expression of α-smooth muscle actin (α-SMA) ( Figure 5 (shown in G), Vimentin ( Figure 5 H) and type I collagen (Col-I) ( Figure 5 α-SMA is an important marker for distinguishing myofibroblasts from fibroblasts, indicating that ColMA-S can significantly promote the differentiation of fibroblasts into myofibroblasts.

[0079] Remarkable cell adhesion and migration as well as superior cell differentiation assays demonstrated that ColMA-S provided a highly bioactive scaffold for HFF-1 cell adhesion, migration and differentiation.

[0080] Example 9 Comparison of the full-thickness skin damage repair effects of low-concentration ColMA scaffolds printed by low-temperature DLP 3D and high-concentration GelMA scaffolds printed by high-temperature DLP 3D

[0081] By establishing a rat full-thickness skin injury model, we evaluated the full-thickness skin injury repair effect of ColMA-S and GelMA-S prepared in Example 5 ( Figure 6As shown in Figure A. Male Sprague-Dawley (SD) rats weighing 150 - 200 g were selected and anesthetized intraperitoneally with 10% sodium pentobarbital at a dose of 0.3 mL / 100 g. The back hair of the rats was shaved and washed with 0.9% sodium chloride injection. The bare skin on the back was disinfected with iodophor. A circular area with a diameter of 1.2 cm was printed on the back using the mouth of a 5 mL centrifuge tube. The upper skin tissue was completely excised with surgical scissors to form a circular full-thickness skin defect wound. DLP 3D printing was used to produce ColMA and GelMA hydrogel scaffolds consistent with the wound model. No material was added to the defect wound in the control group. In the experimental group, ColMA-S and GelMA were applied externally to the defect wound surface. Each group's wound surface was covered with 1 layer of sterilized vaseline gauze and 3 layers of ordinary sterilized gauze to prevent the early shedding of materials and wound infection, and fixed with elastic bandages and zinc oxide PE tape. The growth of SD rats was observed daily, and the wound condition and material shedding were checked. On the 4th, 7th, 14th, 21st, and 28th days after surgery, 6 rats in each group were sacrificed. The tissue at the wound edge and granulation tissue were taken and fixed in 10% formaldehyde fixative for more than 3 days, and then processed by routine paraffin embedding.

[0082] On the 4th, 7th, 14th, 21st, and 28th days respectively, the wound healing conditions of each group were evaluated, and corresponding wound photos were taken. Image J was used to superimpose and visualize the healed skin defect area to more clearly represent the degree of skin repair ( Figure 6 as shown in Figure B). We also used Image J to calculate the wound closure area to determine the wound closure rate ( Figure 6 as shown in Figure C). According to gross observation, the initial wound areas of the control group, GelMA-S group, and ColMA-S group were of similar size; on the 4th day, the wounds of all three groups began to form scabs, wound contraction was visible in the GelMA-S group and ColMA-S group, and no obvious contraction was seen in the control group; on the 7th day, the wound closure rates of the three groups were 35.59%, 60.75%, and 70.56% respectively; on the 14th day, the scabs began to fall off naturally, the wound area of the ColMA-S group was the smallest, and the wound closure rate was 93.33%. The wound closure rates of the control group and GelMA-S group were 65.18% and 81.55% respectively; on the 21st day, the wound of the ColMA-S group completely disappeared and was replaced by scar tissue. There were still wounds in the control group and GelMA-S group, and the wound closure rates were 90.56% and 94.73% respectively; on the 28th day, the scar in the ColMA-S group basically disappeared, and there were still residual wounds in the control group and GelMA-S group. These results indicate that ColMA-S can significantly promote wound healing.

[0083] During the whole experiment, histological changes were regularly observed to evaluate different stages of wound healing. H&E staining was used to evaluate the inflammatory cell infiltration, granulation tissue formation, and epidermal regeneration in the full-thickness skin defect model of SD rats at different time points ( Figure 7as shown in A, C, D, and E). On the 7th day after model establishment, inflammatory cell infiltration occurred in all three groups. Compared with the control group and the GelMA-S group, the level of inflammatory cell infiltration in the ColMA-S group was the lowest, and a large amount of granulation tissue was present. The epidermal regeneration rates of the control group, GelMA-S group, and ColMA-S group were 35.21%, 57.55%, and 71.28% respectively; on the 14th day, the regeneration of granulation tissue in the three groups decreased significantly. The epidermal tissue proliferation in the ColMA-S group was significant, and the epidermal regeneration rate reached 93.06%. The GelMA-S group and the control group were 79.88% and 62.73% respectively; on the 21st day, there was almost no inflammatory cell infiltration in the ColMA-S group, and the regenerated epidermal tissue almost completely covered the wound surface. The epidermal regeneration rate of the ColMA-S group was 98.96%. The epidermal regeneration rates of the control group and the GelMA-S group were 87.74% and 94.70% respectively; on the 28th day, there was still inflammatory cell infiltration in the control group and the GelMA-S group, and the regeneration of granulation tissue was limited. The epidermal regeneration rates were 93.26% and 96.62% respectively. In contrast, the epidermis of the ColMA-S group was completely healed. These experimental results verified that the ColMA scaffold had a significant repair effect, and the newly formed regenerated skin showed characteristics similar to those of normal skin, being dense and orderly.

[0084] The accumulation of collagen fibers in skin regeneration was evaluated by Masson staining ( Figure 7 as shown in B and F). On the 7th day, newly formed collagen fibers appeared at the skin defect sites in the control group, GelMA-S group, and ColMA-S group, with proportions of 12.34%, 22.57%, and 38.06% respectively; on the 14th day, the numbers of collagen fibers in each group were 28.75%, 45.77%, and 55.10% respectively; on the 21st day, the percentages of newly formed collagen fibers in the control group, GelMA-S group, and ColMA-S group were 43.34%, 53.59%, and 72.43% respectively; on the 28th day, a large amount of dense collagen fibers were regenerated in the ColMA-S group, reaching 91.45%, while the collagen fibers in the control group and the GelMA-S group were sparsely distributed, being 55.78% and 66.62% respectively. The results of Masson staining showed that the formation of collagen fibers in the control group and the GelMA-S group was disordered, while the collagen deposition in the ColMA-S group was obvious and the collagen fibers were arranged regularly. These results all indicated that ColMA-S could significantly promote the effective repair of full-thickness skin defect wounds.

Claims

1. A method for low-concentration collagen low-temperature DLP 3D printing of ColMA biomaterials, characterized in that, the printing method is: dissolving photosensitive collagen ColMA in acetic acid solution and then mixing it evenly with the photoinitiator LAP solution, and performing DLP 3D printing at 4-25 °C under blue light irradiation with a wavelength of 405 nm to obtain a collagen biomaterial; the ColMA is collagen modified with methacrylic anhydride.

2. The method according to claim 1, characterized in that, the method comprises the following steps: (1) Obtaining photosensitive collagen ColMA by methacrylation modification of collagen; (2) Dissolving the photosensitive collagen ColMA obtained in step (1) in acetic acid solution to obtain a photosensitive collagen ColMA solution; (3) Adding a photoinitiator LAP to the photosensitive collagen ColMA solution described in step (2) to prepare a collagen bioink; (4) Adding the bioink obtained in step (3) to a DLP 3D printer with blue light of a wavelength of 405 nm, and performing layer-by-layer printing by photocuring molding at 4-25 °C, with a curing time of 1-5 s for each layer; (5) Secondarily curing the material printed and formed in step (4) under blue light with a wavelength of 405 nm at 4-25 °C, with a curing time of 1-8 s, to obtain a DLP 3D printed ColMA biomaterial.

3. The method according to claim 2, characterized in that, the pH of the ColMA solution is 2.5-3.

0.

4. The method according to claim 2, characterized in that, the concentration ratio of the ColMA solution to the LAP solution is 1:0.1-5.

5. The method according to claim 2, characterized in that, the concentration of the ColMA solution is 0.1-1% m / v.

6. The method according to claim 2, characterized in that, the final concentration of the LAP solution is 0.1-0.5% m / v.

7. The method according to claim 2, characterized in that, the preparation method of the photosensitive collagen ColMA is: dissolving collagen in acetic acid solution, adjusting the pH of the reaction system to 7-9, adding an excessive amount of methacrylic anhydride, reacting for 24 hrs under ice bath conditions, dialyzing and purifying, and freeze-drying to obtain the photosensitive collagen ColMA.

8. The method according to claim 7, characterized in that, the concentration of the collagen stock solution is 0.01-1% m / v.

9. The method according to claim 7, characterized in that, the molar ratio of collagen to methacrylic anhydride is 1:60-100.

10. Application of the ColMA biomaterial prepared by any one of the methods according to claims 1-9 in the preparation of skin injury repair products.

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