Hybrid Tissue Engineered Constructs

Abstract

The present invention provides a method for producing a tissue engineered construct. The tissue engineered construct of the present disclosure is characterized by a surface treatment of a scaffold to modify the surface properties of the scaffold. A hydrophilic hydrogel network is physically crosslinked to the treated surface via a charged polymer and salt ions. Biological materials are entrapped in the physically crosslinked hydrogel network and thereby loaded onto the hydrogel network. Covalently reactive macromonomers are chemically crosslinked to the hydrophilic hydrogel network to enhance the physical crosslinking of the hydrophilic hydrogel network to itself and to the scaffold. The tissue engineered construct of the present disclosure allows for the delivery of therapeutic agents including cells and / or biomolecules and has structural support and a defined shape for regenerative medicine applications.

Description

[Technical field]

[0001] (Related Applications) This application claims priority to U.S. Provisional Patent Application No. 63 / 289,431, filed December 14, 2021, U.S. Provisional Patent Application No. 63 / 304,216, filed January 28, 2022, U.S. Provisional Patent Application No. 63 / 289,447, filed December 14, 2021, and U.S. Provisional Patent Application No. 63 / 304,207, filed January 28, 2022. The entire disclosures of each of the above applications are incorporated herein by reference.

[0002] (Technical field) The present invention relates to tissue engineered constructs and methods for producing same. [Background technology]

[0003] Reconstruction of large bone defects resulting from bone injury, bone disease, or congenital disorders remains a significant clinical problem. More than 500,000 bone grafting procedures are performed annually in the United States and more than 2 million bone grafting procedures are performed annually worldwide. Autogenous bone is the gold standard for bone grafting. However, additional surgery is required to harvest autogenous bone from the donor site, the amount of bone harvested is limited for reconstruction of large bone defects, and the donor site may become diseased. Therefore, allografts have been widely used for bone reconstruction as an alternative to autografts. However, the long-term failure rate of allografts in the treatment of large severe bone defects is 25%-60% due to various complications. In addition, according to the CDC (Centers for Disease Control and Prevention), the use of frozen allografts carries a potential risk of disease transmission. Although commercially available demineralized bone matrix (DBM) contains osteoinductive factors, DBM alone does not provide the structural and mechanical support required for reconstruction of large bone defects. Therefore, synthetic scaffolds as bone graft substitutes have attracted attention in recent years.

[0004] Scaffolds for bone tissue engineering should ideally be biocompatible, bioresorbable, mechanically stable, porous, osteoconductive, and osteoinductive. Biocompatible and bioresorbable FDA-approved polyesters, such as polycaprolactone (PCL), polyglycolic acid (PGA), polylactic acid (PLA), and their copolymers (e.g., polylactic-co-glycolic acid (PLGA)), are the most widely used synthetic polymers in bone tissue engineering. Various methods have been used to fabricate porous polyester-based scaffolds, including molding, solvent casting / porogen leaching, gas foaming, laser drilling, and 3D printing. Among these techniques, 3D printing can precisely control the structure and porosity of the scaffold. Proper control of the porosity of 3D-printed scaffolds is particularly important for bone tissue engineering, since the presence of interconnected pores with pore sizes of 300 μm or more is essential for cell migration and intrapore bone formation.

[0005] The osteoconductivity and mechanical properties of polyester-based scaffolds are enhanced by incorporating calcium phosphate bioceramics. For example, osteogenic differentiation of mouse preosteoblasts (MC3T3-E1) on 3D-printed PCL / β-tricalcium phosphate (TCP) substrates was significantly higher than that on pristine PCL substrates. We previously demonstrated that the Young's modulus of 3D-printed PCL-TCP scaffolds can be tuned in the range of 12–188 MPa by varying the TCP content and scaffold porosity. Furthermore, clinically available electron beam sterilization did not adversely affect the mechanical and bioactive properties of PCL-TCP scaffolds.

[0006] 3D-printed polymer / ceramic scaffolds are biocompatible, bioresorbable, mechanically stable, porous, and osteoconductive, but lack osteoinductive factors that induce osteogenic differentiation and promote bone healing. Therefore, it is necessary to incorporate osteoinductive proteins into 3D-printed scaffolds, especially for treating large bone defects. Surface coatings have been used to immobilize proteins on the surface of 3D-printed scaffolds for tissue engineering applications. However, protein loading on thin coatings is generally limited and the release rate is fast. For example, the loading amount of BSA on 3D-printed hydroxyapatite scaffolds coated with chitosan and sodium hyaluronate by layer-by-layer (LBL) technique was lower and the release rate was faster than that on uncoated scaffolds.

[0007] Hydrogels have been used for protein delivery and cell encapsulation in tissue engineering applications. 3D-printed polymeric networks of hydrogels with high water content provide a platform for sufficient protein loading and sustained release. However, hydrogels have poor stiffness and structural integrity, limiting their use as stand-alone 3D-printed scaffolds. In addition, incorporating soft hydrogels into rigid 3D-printed scaffolds is difficult due to the mismatch in mechanical properties at the interface. Therefore, methods have been used to incorporate hydrogels into rigid scaffolds by filling the 3D-printed porous scaffolds with hydrogel precursor solutions and then gelling the hydrogel with light or heat. For example, a surface tension-assisted method is used to fill the pores of the 3D-printed constructs with photocrosslinkable methacrylated gelatin hydrogel. Multi-material 3D printing methods have been used to fabricate porous polymer / hydrogel composite scaffolds.

[0008] Maintaining the porous structure of the scaffold after incorporating hydrogel is essential for cell migration, tissue integration, and vascularization for the diffusion of nutrients and oxygen in tissue engineering applications. A tissue-organ integration printer was used to sequentially print gelatin / fibrinogen-based hydrogels along with PCL structural supports. Stanford University developed a 3D hybrid bioprinting technology (Hybprinter) and used it to fabricate composite scaffolds from PCL and polyethylene glycol diacrylate (PEGDA) hydrogels. Despite its technological importance, multi-material printing requires long manufacturing times and special expensive 3D printers due to multiple iterations between materials during printing. Moreover, the simultaneous printing of polymers / ceramics and hydrogels hampers the surface treatment of the scaffolds and the improvement of the integration of soft and hard materials at the interface. Summary of the Invention [Means for solving the problem]

[0009] The present invention, in one example, provides porous or non-porous biologically loaded multi-material constructs (hereinafter referred to as Hybrid Tissue Engineered Constructs (HyTECs)) for applications in regenerative medicine and disease treatment.

[0010] Constructs and devices made of polymers, ceramics, metals, or composites in porous or non-porous form are widely used as implants in regenerative medicine. Various techniques, such as 3D printing and casting, have been used to fabricate porous implants. Coating techniques, such as layer-by-layer coating and adhesive coating, have also been used to load biomolecules on the surface of porous implants. However, these coating techniques can only load limited amounts of biomolecules. Loading implants with large or tunable amounts of biomolecules is particularly important because the effective amount of biomolecules is often high in vivo and varies for various indications.

[0011] Biological materials (biomolecules, drugs, and / or cells) can be loaded onto the implant by filling the porous structure of the implant with a hydrogel loaded with the biological material. However, filling the porous space of the implant with a hydrogel blocks the pores, inhibiting or preventing cell recruitment, cell migration, vascular invasion, tissue regeneration, and integration with the surrounding tissue.

[0012] To address at least this issue, the inventors of the present invention have developed a method for fabricating HyTEC that allows the incorporation of biological materials on the surface of a porous scaffold while keeping the interconnected pores unblocked, or on the surface of a non-porous implant through a uniformly thick hydrogel layer (FIG. 1). For proof-of-concept, the inventors loaded model proteins and cells onto the surfaces of 3D-printed biodegradable polycaprolactone and β-tricalcium phosphate (PCL-TCP) as model polymeric ceramic porous scaffolds, PCL-TCP rods as model polymeric ceramic non-porous implants, and stainless steel needles as model metals.

[0013] The surface of a porous or non-porous scaffold is treated in three consecutive steps (1) to (3) as follows (Figures 2A and 2B). (1) Increase hydrophilicity / reactivity / roughness. (2) Improve the adhesiveness of hydrogels. (3) Promotes (induces) surface-initiated crosslinking.

[0014] The surface of the scaffold was loaded with a hydrogel layer by surface-initiated physical crosslinking followed by covalent crosslinking.

[0015] Hydrophilicity / reactivity / roughness enhancement

[0016] Sodium hydroxide (NaOH) treatment and freezing / thawing were used to increase the hydrophilicity / reactivity of the surface of the PCL-TCP scaffolds. Other treatment methods, such as plasma treatment and acid treatment, may also be used to increase the hydrophilicity / reactivity / roughness of the scaffold surface.

[0017] Improved hydrogel adhesion

[0018] To improve the adhesion of hydrogels, the surface of the scaffolds was coated with molecules that have functional groups that can be covalently bonded. For example, reactive aminopropylmethacrylamide (APMA) and gelatin methacrylate (GelMA) were coupled to the surface of PCL-TCP scaffolds using carbodiimide chemistry (Figure 2A, B). Amine-reactive (N-hydroxysuccinimide) ester diazine (NHS-diazirine, succinimidyl 4,4´-adipentanoate) was coupled to the surface of stainless steel.

[0019] Accelerating surface-initiated crosslinking

[0020] To promote (induce) surface-initiated physical crosslinking, calcium chloride (CaCl 2 ) or calcium sulfate (CaSO 4 ) is deposited on the surface of the implant. Other salts of divalent cations (e.g., Ca 2+ , Mg 2+ , Sr 2+ ), or other salts of polyvalent cations (e.g., Ti 4+ , Al 3+ ) may be used for surface initiated physical crosslinking.

[0021] After these three steps of surface treatment, the scaffolds were immersed in a hydrogel precursor solution containing alginate, covalently reactive macromonomers, initiators, and biological materials (biomolecules, drugs, and / or cells). Polyethylene glycol dimethacrylate (PEGDMA) and GelMA were used as covalently reactive macromonomers (Figure 2A, Figure 2B). Other macromonomers and crosslinkers with double bonds or other covalently reactive functional groups (e.g., NHS groups for amine reactions or SH groups for Michael addition) may also be used in the hydrogel precursor solution.

[0022] When the surface-treated scaffold was immersed in a hydrogel precursor solution, calcium ions diffused from the surface of the scaffold into the solution and crosslinked the alginate near the surface, forming a hydrogel layer on the surface of the scaffold. The macromonomers in the physically crosslinked hydrogel were then covalently crosslinked in the next step to form a robust interpenetrating network. Porous and nonporous HyTEC were fabricated using chemical initiators (APS / TEMED) and photoinitiators (lithium phenyl-2,4,6-trimethylbenzoylphosphinate) (Figures 2A and 2B). Other chemical initiators, visible light initiators, UV initiators, or thermal initiators may be used to initiate the covalent crosslinking reaction.

[0023] adjustment

[0024] The hydrogel loading rate and hydrogel thickness can be adjusted by changing the process parameters. For example, the thickness of the hydrogel layer on the porous PCL-TCP scaffold can be adjusted by changing the NaOH surface treatment time and the CaCl content in the solution used for calcium deposition. 2 The thickness of the coating on the construct can be spatially tuned / controlled from 0 to high by immersing different parts in different solutions.

[0025] result

[0026] 3D-printed porous PCL-TCP scaffolds with various porosities maintained their porosity after hydrogel loading (Figure 3A,B). When BMP2 (as a model protein) was loaded onto HyTECs, BMP2 was released over 28 days in vitro. The presence of a thick hydrogel layer allowed a wide range of biomolecule doses to be loaded onto HyTECs. Biomolecule-loaded HyTECs could be lyophilized, stored, and terminally sterilized using a fractionated electron beam (EB) method (Figure 4A-4C). Live cells could be encapsulated in the hydrogel layer of HyTECs (Figure 5A-5D). Furthermore, hydrogels were loaded onto stainless steel needles as a proof-of-concept for the fabrication of metal-based HyTECs (Figure 6).

[0027] Depending on the application, HyTEC may contain cells or may be cell-free, with or without therapeutic biomolecules. HyTEC may be raw, frozen, or lyophilized. Also, the scaffold may be porous or non-porous and may be made from polymers (e.g., polyesters), metals, ceramics, or composites. The materials used for hydrogel macromonomers, gelation initiators, salts for calcium deposition, and surface treatments may vary depending on the scaffold surface chemistry and application.

[0028] Purpose

[0029] The embodiments of the present invention can be applied or used as follows, without being limited in scope of the present invention: HyTEC is used to deliver therapeutic agents including cells and / or biomolecules, and has structural support and defined geometry for regenerative medicine applications. Examples are given below. · Osteoconductive 3D printed constructs for delivery of osteoconductive proteins and bone forming cells, and treatment of bone defects. 3D printed constructs for protein and / or cell delivery and treatment of soft tissue defects. Antibiotic delivery, 3D printed scaffolds. Painkiller delivery, 3D printed scaffolds. Protein and drug delivery, metallic implants. Delivery of angiogenic proteins or cells to induce angiogenesis in regenerative medicine. Localized delivery of therapeutic agents to cancer tissue. ·Local delivery of beta cells for insulin secretion in diabetic patients.

[0030] advantage

[0031] Embodiments of the present invention have advantages over existing approaches and constructs. In contrast to methods based on thin coatings (e.g., layer-by-layer coating or coating constructs with absorbents), the HyTEC technology allows porous (or non-porous) constructs to be loaded with higher doses, broader doses, or a variety of therapeutic or biological agents. Additionally, cells can be encapsulated in HyTEC in contrast to constructs with thin coatings.

[0032] The advantages of HyTEC technology over multi-material printing are: despite its technological importance, multi-material printing requires long production times and special expensive 3D printers due to multiple iterations between materials during printing, and the nature of the various printing mechanisms limits the choice of processing parameters. Moreover, the simultaneous printing of polymers / ceramics and hydrogels hampers the surface treatment of the scaffold and the improvement of the integration of soft and hard materials at the interface.

[0033] In one embodiment, bioactive implants (e.g., HyTEC constructs) can be coated with absorbable polyesters (e.g., PCL, PLA, PLGA) or other absorbable polymers (e.g., polyurethane) to slow the release of therapeutic agents. HyTEC is an abbreviation for hybrid tissue engineered constructs, and is a bioactive implant. A schematic diagram of the method used to coat HyTEC constructs is shown in Figure 7. For example, protein-loaded HyTEC can be frozen overnight at -80°C, then frozen at -20°C for 10 minutes, and then immersed in a PCL solution in acetone or chloroform (2%-20%) to deposit a PCL layer on the HyTEC, creating modified HyTEC (mHyTEC). The PCL-coated HyTEC is then air-dried at 0-4°C. The physical properties of the mHyTEC construct and the release rate of the protein can be adjusted by varying the concentration of the PCL solution and the number of PCL layers deposited. Representative images of mHyTECs with one layer of PCL coating (mHyTEC(1L)) or three layers of PCL coating (mHyTEC(3L)) are shown in Figure 8. After 14 days, 92% of the encapsulated BSA was released from uncoated BSA-loaded HyTECs, but when a PCL / acetone (10% wt / v) solution was used to form one or three protective PCL coatings, the BSA release rate decreased to 92% after 70 days or 80% after 91 days (Figure 9). Also, the amount of bone morphogenetic protein 2 (BMP2) protein released from BMP2-loaded HyTEC constructs after 28 days in PBS decreased from 84% to 62% or 24% when one or three layers of PCL coatings were deposited (Figure 10).

[0034] In another example, the present invention provides a method for producing a tissue engineered construct. A scaffold having a surface with predetermined surface properties is provided. The surface of the scaffold is treated to modify the surface properties of the scaffold. Optionally, the surface of the scaffold is coated with a molecule capable of covalent bonding to promote chemical crosslinking to the surface of the scaffold. Surface initiated physical crosslinking is promoted by depositing salts on the surface of the scaffold or on the surface that is optionally coated with the molecule capable of covalent bonding. A hydrogel precursor solution is provided that contains a charged polymer, a covalently reactive macromonomer, an initiator, and a biological material. A physically crosslinked hydrophilic hydrogel network is formed on the surface of the scaffold by immersing the scaffold in the hydrogel precursor solution. The formation of the hydrophilic hydrogel network is controlled by the release of salt ions from the surface of the scaffold and the physical crosslinking of the salt ions released from the surface of the scaffold with the charged polymer. The biological material is entrapped in the hydrophilic hydrogel network during the formation of the hydrophilic hydrogel network, thereby being carried by the hydrogel network. The scaffold with the formed hydrophilic hydrogel network is removed from the hydrogel precursor solution. The hydrophilic hydrogel network is chemically crosslinked with covalently reactive macromonomers to enhance the physical crosslinking of the hydrophilic hydrogel network to itself and to the scaffold. Optionally (following optional surface treatment), a covalently bondable molecule coated on the surface of the scaffold is chemically crosslinked with the covalently reactive macromonomers to enhance adhesion of the chemically and physically crosslinked hydrophilic hydrogel network to the scaffold.

[0035] In a further step, the tissue engineered construct is optionally frozen or lyophilized.

[0036] In a further step, the surface of the scaffold is further treated to increase the hydrophilicity and / or roughness of the surface of the scaffold.

[0037] The scaffold has an interconnected porous structure, and the above steps of the method are controlled such that the hydrophilic hydrogel network is physically and chemically crosslinked and chemically bonded to the interconnected porous structure of the scaffold, and the pores of the interconnected porous structure are maintained by the above steps of the method such that they are capable of accommodating biological material.

[0038] In yet another step, the tissue engineered construct is coated with one or more coating layers.

[0039] In yet another step, the scaffold has an interconnected porous structure and the coating controls the pore size of the interconnected porous structure.

[0040] In another embodiment, the present invention provides a tissue engineered construct. The tissue engineered construct of the present disclosure includes a scaffold having a surface treated to modify the surface properties. A hydrophilic hydrogel network is physically crosslinked to the treated surface via a charged polymer and salt ions. Biological material is entrapped in the physically crosslinked hydrogel network, thereby being supported by the hydrogel network. Covalently bonded reactive macromonomers are chemically crosslinked to the hydrophilic hydrogel network to enhance the physical crosslinking of the hydrophilic hydrogel network to itself and to the scaffold.

[0041] In a variation of the tissue engineering of the present disclosure, the surface of the scaffold is coated with a molecule capable of covalent bonding and a covalent reactive macromonomer is chemically crosslinked to the molecule capable of covalent bonding to enhance adhesion of the chemically and physically crosslinked hydrophilic hydrogel network to the scaffold. The tissue engineering of the present disclosure may have one or more coating layers. The scaffold is an interconnected porous scaffold, and the biological material is carried in the pores of the interconnected porous scaffold.

[0042] Fused deposition modeling is an effective method to print three-dimensional (3D) bioresorbable scaffolds and medical devices with well-controlled porosity, internal microstructure, and overall shape for biomedical applications. However, proteins and living cells cannot survive hot extrusion. As a further feature of the present invention, the disclosed hybrid tissue engineered constructs (HyTECs) are designed to allow the incorporation of biological materials (e.g., proteins and cells) through a uniform thick hydrogel layer on the surface of 3D-printed porous scaffolds while keeping the interconnected pores unblocked, or on the surface of non-porous implants. 3D-printed biodegradable polycaprolactone-β-tricalcium phosphate (PCL-TCP) was used as a model porous scaffold, PCL-TCP rods were used as a model non-porous implant, and bone morphogenetic protein-2 (BMP-2) was used as a model protein for bone tissue engineering applications. To increase hydrophilicity, improve hydrogel adhesion, and promote surface-initiated crosslinking, the surface of the PCL-TCP construct was treated in three successive steps. A hydrogel layer was loaded onto the surface of the scaffold by surface-initiated physical crosslinking followed by covalent crosslinking. As a result, the surface treatment of the scaffold improved the adhesion of the hydrogel to the surface of the scaffold without adversely affecting the mechanical and surface properties of the scaffold. By adjusting the surface treatment parameters, the average thickness of the hydrogel layer loaded onto the surface of the scaffold could be controlled in the range of 100-600 μm. In addition, the 3D-printed scaffolds with porosity of 50-80% maintained a pore size of 140-1100 μm after loading of the hydrogel onto the surface of the scaffold. Two cell types (hMSCs and C 3 H 10Cell viability and proliferation tests using EB sterilization showed that hydrogel loading did not adversely affect the biocompatibility of the scaffolds. BMP-2-loaded hydrogel scaffolds sustainedly released BMP-2 over a period of 35 days. Freeze-drying and e-beam sterilization of hydrogel-loaded PCL-TCP scaffolds did not adversely affect the mechanical properties of the scaffolds, but did adversely affect the amount of active BMP-2 released from the scaffolds. The amount of active BMP-2 released from sterilized freeze-dried HyTEC constructs was improved two-fold by using a fractionated e-beam irradiation strategy. The thick hydrogel layer enabled the HyTEC constructs to load encapsulated live cells with over 92% cell viability after 7 days. In summary, the disclosed HyTEC fabrication method holds great promise for improving the payload capacity of biological materials in porous (or non-porous) polyester-based 3D printed tissue engineering scaffolds while maintaining unobstructed interconnected pores for improved tissue integration and survival. [Brief description of the drawings]

[0043] If necessary, for further interpretation of the greyscales in the figures, the reader is referred to the priority documents for each figure.

[0044] [Figure 1] FIG. 1 is a schematic diagram of a HyTEC, in accordance with an exemplary embodiment of the present invention. [Figure 2A]FIG. 2A is a schematic diagram of a process for fabricating a porous HyTEC according to an exemplary embodiment of the present invention. FIG. 2A is a schematic diagram of a process for fabricating a 3D printed porous PCL-TCP / hydrogel HyTEC construct. After 3D printing, the PCL-TCP scaffold was treated with NaOH to increase the hydrophilicity of the surface by cleaving the ester bonds to the carboxyl and hydroxyl groups. Then, reactive double bonds were introduced on the surface of the scaffold by grafting APMA to the carboxyl groups using carbodiimide chemistry. The scaffold was then immersed in a CaCl2 solution and then dried under vacuum. When the CaCl2 surface-treated scaffold was immersed in a hydrogel precursor solution, the CaCl2 deposited on the surface of the scaffold diffused from the surface of the scaffold into the solution and crosslinked the alginic acid near the surface, thereby forming a hydrogel layer on the surface of the scaffold. The PEGDMA macromonomers in the physically crosslinked hydrogels were covalently crosslinked in the next step to form a robust interpenetrating network. The hydrogel network was attached to the scaffold surface via the reaction of the double bonds of the PEGDMA macromonomers with the double bonds of the APMA grafted onto the scaffold surface. [Figure 2B] FIG. 2B is a schematic diagram of a process for fabricating non-porous HyTEC, according to an exemplary embodiment of the invention. [Figure 3A] FIG. 3A shows porous HyTEC having various porosities according to an exemplary embodiment of the present invention. [Figure 3B] FIG. 3B illustrates the effect of HyTEC porosity on supported hydrogel thickness, according to an exemplary embodiment of the present invention. [Figure 4A] FIG. 4A shows the release kinetics of BMP2 from lyophilized HyTEC, according to an exemplary embodiment of the present invention. [Figure 4B] Figure 4B shows the effect of storing lyophilized HyTECs at 4 °C for 2 months on the release kinetics of BMP2. [Figure 4C]FIG. 4C shows the effect of electron beam sterilization with fractionated irradiation on the release kinetics of BMP2 from lyophilized HyTEC. [Figure 5A] FIG. 5A shows an image of a non-porous HyTEC containing human mesenchymal stem cells, according to an exemplary embodiment of the invention. [Figure 5B] FIG. 5B shows live (green) and dead (red) cells encapsulated within non-porous HyTEC. [Figure 5C] FIG. 5C shows the viability of human mesenchymal stem cells (hMSCs) in HyTECs over a 7-day period. [Figure 5D] FIG. 5D shows the DNA content of HyTEC containing cells over a 14 day period. [Figure 6] FIG. 6 shows a model metal-based HyTEC fabricated on a stainless steel needle, according to an exemplary embodiment of the present invention. [Figure 7] FIG. 7 shows a schematic diagram of the method used to coat a bioactive implant (HyTEC construct), according to an exemplary embodiment of the present invention. [Figure 8] FIG. 8 shows images of mHyTEC with one layer of PCL coating (mHyTEC(1L)) and three layers of PCL coating (mHyTEC(3L)) fabricated using a PCL / acetone (10% wt / v) solution according to an exemplary embodiment of the invention. [Figure 9] FIG. 9 shows the release kinetics of BSA from HyTEC constructs without a PCL protective coating (control), with one layer of a PCL coating made using a PCL / chloroform (10% wt / v) solution (PCL / chloroform: 1 L), with one layer of a PCL coating made using a PCL / acetone (10% wt / v) solution (PCL / acetone: 1 L), with three layers of a PCL coating made using a PCL / chloroform (10% wt / v) solution (PCL / chloroform: 3 L), and with three layers of a PCL coating made using a PCL / acetone (10% wt / v) solution (PCL / acetone: 3 L) in accordance with an exemplary embodiment of the invention. [Figure 10] FIG. 10 shows the release kinetics of rhBMP2 from HyTEC constructs (HyTEC) without a PCL protective coating, with one layer of PCL coating made using a PCL / acetone (10% wt / v) solution (mHyTEC(1L)), and with three layers of PCL coating made using a PCL / acetone (10% wt / v) solution (mHyTEC(3L)) according to an exemplary embodiment of the present invention. [Figure 11A] FIG. 11A shows the effect of APMA concentration in a reaction solution on the density of grafted APMA on the scaffold surface, according to an exemplary embodiment of the present invention. [Figure 11B] FIG. 11B shows the effect of surface modification with APMA on the contact angle of PCL-TCP structures. [Figure 11C] Figure 11C shows PCL-TCP scaffolds with porosity ranging from 0% to 80% fabricated by 3D printing, which were treated with NaOH and APMA (2.5 mg / mL) and used to evaluate the effect of surface treatment on the mechanical properties of the scaffolds. [Figure 11D] Figure 11D shows the Young's modulus of PCL-TCP scaffolds without surface treatment (Untreated; B) and with NaOH / APMA treatment (Treated / A-2.5; R). Error bars represent the mean ± 1 SD (n=3). [Figure 11E] Figure 11E shows the yield stress of PCL-TCP scaffolds without surface treatment (Untreated; B) and with NaOH / APMA treatment (Treated / A-2.5; R). Error bars represent the mean ± 1 SD (n=3). [Figure 12A] FIG. 12A shows the effect of NaOH treatment time on hydrogel coating in PCL-TCP scaffolds with 80% porosity, according to an exemplary embodiment of the present invention. [Figure 12B] FIG. 12B shows the effect of NaOH treatment time on the release of calcium ions from the surface of the scaffold. [Figure 12C]FIG. 12C shows the effect of NaOH treatment time on hydrogel thickness. [Figure 12D] Figure 12D shows the effect of NaOH treatment time on the percentage of filled pores. In Figures 12A, 12C, and 12D, the CaCl2 concentration in the treatment solution was 100 mg / mL, and the PEGDMA concentration in the hydrogel precursor solution was 20% (wt / vol). Error bars represent the mean ± 1 SD. [Figure 13A] FIG. 13A shows the effect of CaCl2 concentration in treatment solutions on hydrogel coating on PCL-TCP scaffolds with 80% porosity, according to an exemplary embodiment of the present invention. [Figure 13B] FIG. 13B shows the effect of the concentration of CaCl2 in the treatment solution on the release of calcium ions from the surface of the scaffold. [Figure 13C] FIG. 13C shows the effect of the concentration of CaCl in the treatment solution on the thickness of the hydrogel. [Figure 13D] Figure 13D shows the effect of CaCl2 concentration in the treatment solution on the percentage of filled pores. In Figures 13A, 13C, and 13D, the NaOH treatment time was 60 min, and the PEGDMA concentration in the hydrogel precursor solution was 20% (wt / vol). Error bars represent the mean ± 1 SD. [Figure 14A] FIG. 14A illustrates the effect of scaffold porosity on a hydrogel coating on an interconnected porous scaffold, according to an exemplary embodiment of the invention. [Figure 14B] FIG. 14B shows the effect of scaffold porosity on hydrogel thickness. [Figure 14C] FIG. 14C shows the effect of scaffold porosity on the percentage of filled pores. [Figure 14D] FIG. 14D shows the effect of scaffold porosity on the release of calcium ions from the surface of the scaffold. [Figure 14E]FIG. 14E shows the effect of scaffold porosity on the release of calcium ions per unit weight of scaffold. [Figure 14F] Figure 14F shows the effect of scaffold porosity on hydrogel loading. In Figures 14A-14F, the NaOH treatment time was 60 min, the concentration of CaCl2 in the treatment solution was 100 mg / mL, and the concentration of PEGDMA in the hydrogel precursor solution was 20% (wt / vol). Error bars represent the mean ± 1 SD. [Figure 15A] 15A-15G show the effect of freeze-drying on the properties of PCL-TCP scaffolds and HyTEC according to an exemplary embodiment of the present invention: Figure 15A shows the effect of freeze-drying on the Young's modulus of PCL-TCP scaffolds at various porosities. [Figure 15B] FIG. 15B shows the effect of freeze-drying on the yield stress of PCL-TCP scaffolds at various porosities. [Figure 15C] FIG. 15C shows the effect of freeze-drying and rehydration on a hydrogel coating of HyTEC with 80% porosity. [Figure 15D] FIG. 15D shows the effect of NaOH treatment time on the hydrogel thickness of filled pores in HyTEC with 80% porosity before lyophilization and after lyophilization and rehydration. [Figure 15E] FIG. 15E shows the effect of NaOH treatment time on the percentage of filled pores in HyTEC with 80% porosity before lyophilization and after lyophilization and rehydration. [Figure 15F] FIG. 15F shows the effect of CaCl concentration on hydrogel thickness in HyTEC with 80% porosity before lyophilization and after lyophilization and rehydration. [Figure 15G]Figure 15G shows the effect of CaCl2 concentration on the percentage of filled pores in HyTEC with 80% porosity before freeze-drying and after freeze-drying and rehydration. In Figure 15D-E, the CaCl2 concentration in the treatment solution was 100 mg / mL, and the PEGDMA concentration in the hydrogel precursor solution was 20% (wt / vol). Also, in Figure 15F-G, the NaOH treatment time was 60 min, and the PEGDMA concentration in the hydrogel precursor solution was 20% (wt / vol). "Asterisks" represent statistically significant differences between non-freeze-dried and freeze-dried / rehydrated samples. Error bars correspond to the mean ± 1 SD. B represents before freeze-drying, and F represents after freeze-drying. [Figure 16A] 16A-H show characterization of mechanical properties and BMP-2 release kinetics of HyTEC under various conditions according to an exemplary embodiment of the present invention. [Figure 16B] 16A-B show the structure of the 3D printed PCL-TCP device used to measure the adhesion of the hydrogel to the scaffold. The hydrogel is embedded within the gap between two concentric cylinders. [Figure 16C] The hydrogel-integrated device was then placed on an Instron instrument, the two bridges connecting the inner and outer cylinders were cut, and the interfacial stiffness was measured by a push-out test (Figure 16C). [Figure 16D] FIG. 16D shows the effect of APMA surface treatment on the interfacial stiffness of the hydrogel. [Figure 16E] FIG. 16E shows the effect of PEGDMA polymer concentration in the hydrogel precursor solution on the interfacial stiffness of the hydrogel. [Figure 16F] FIG. 16F shows the effect of PEGDMA polymer concentration in the hydrogel precursor solution on hydrogel loading in PCL-TCP scaffolds with 80% porosity. [Figure 16G] FIG. 16G shows the effect of PEGDMA polymer concentration in the hydrogel precursor solution on BMP2 release kinetics in PCL-TCP scaffolds with 80% porosity. [Figure 16H]Figure 16H shows the release kinetics of BMP2 from HyTEC with 80% porosity, containing untreated BMP2 (red), freeze-dried BMP2 (green), and freeze-dried and e-beam sterilized BMP2 (blue). In Figure 16F-H, the NaOH treatment time was 60 min and the CaCl2 concentration was 100 mg / mL. Error bars correspond to the mean ± 1 SD. [Figure 17A] 17A-H show cell viability, proliferation, and osteogenic differentiation of HyTEC in vitro according to an exemplary embodiment of the present invention: Figure 17A shows the effect of hydrogel loading on normalized viability of hMSC and C3H10 cultured in preconditioned DMEM medium. [Figure 17B] FIG. 17B shows the effect of hydrogel loading rate on hMSC cells cultured in preconditioned DMEM medium. [Figure 17C] FIG. 17C shows the effect of hydrogel loading rate on C3H10 cells cultured in preconditioned DMEM medium. [Figure 17D] FIG. 17D shows the ALP activity of hMSC cells cultured in DMEM medium (C), DMEM medium pretreated with HyTEC without BMP2 (scaffold+gel; P), DMEM medium pretreated with HyTEC containing BMP2 (scaffold+gel / BMP2; R), DMEM medium supplemented with 1.5 μg / mL BMP2 for 3 days (BMP2 in medium (3d); B), and DMEM medium supplemented with 214 ng / mL BMP2 for 21 days (BMP2 in medium (21d); G). [Figure 17E]FIG. 17E shows ALP activity of C3H10 cells cultured in DMEM medium (C), HyTEC-pretreated DMEM medium without BMP2 (scaffold+gel; P), HyTEC-pretreated DMEM medium with BMP2 (scaffold+gel / BMP2; R), DMEM medium supplemented with 1.5 μg / mL BMP2 for 3 days (BMP2 in medium (3d); B), and DMEM medium supplemented with 214 ng / mL BMP2 for 21 days (BMP2 in medium (21d); G). "Asterisks" represent statistically significant differences between non-lyophilized and lyophilized / rehydrated samples. Error bars correspond to the mean ± 1 SD. [Figure 18A] 18A-I show the fabrication process and characterization of a representative non-porous HyTEC according to an exemplary embodiment of the present invention. Figure 18A shows a schematic of the procedure for coating a non-porous PCL-TCP rod with a bioabsorbable hydrogel. [Figure 18B] FIG. 18B shows an SEM image of the surface of the hydrogel supported on the non-porous PCL-TCP rod. [Figure 18C] FIG. 18C shows an SEM image of the surface of the hydrogel supported on the non-porous PCL-TCP rod. [Figure 18D] FIG. 18D shows the average loading of hydrogels without BMP2 (G) and with BMP2 (R) on treated and calcium-deposited PCL-TCP rods. [Figure 18E] Figure 18E shows the release kinetics of BMP2 from freeze-dried BMP2-containing hydrogel-loaded PCL-TCP rods without electron beam sterilization (G), after sterilization by a single dose of electron beam irradiation (B), and after sterilization by split doses of electron beam irradiation (R). [Figure 18F] FIG. 18F shows an image of cell-loaded non-porous HyTEC. [Figure 18G] FIG. 18G shows an image of non-porous HyTEC with stained live (G) / dead (R) cells. [Figure 18H] FIG. 18H shows the viability of hMSCs encapsulated in non-porous HyTEC. [Figure 18I]Figure 18I shows the DNA content of cell-loaded non-porous HyTECs. Error bars represent the mean ± 1 SD. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0045] definition

[0046] The following detailed description is an exemplary embodiment of a method for producing / manufacturing a tissue engineered construct, and structural features of the tissue engineered construct. In general, the following definitions of terms can be used as a guide within the scope of the present invention. A scaffold is defined as a porous or non-porous three-dimensional construct made from polymers, ceramics, metals, or composites. Surface treatments include base (e.g., NaOH) treatment, acid treatment, plasma treatment, freeze / thaw, and other methods. Surfaces treated to modify surface properties are defined as surfaces that have had their roughness increased by chemical or physical treatment (e.g., base, acid, plasma, freeze / thaw treatments, etc.). Covalently bondable molecules that facilitate chemical cross-linking to surfaces are molecules that have a functional group for bonding to the surface and another functional group that can bond to other molecules. Examples include aminopropylmethacrylamide (APMA), gelatin methacrylate (GelMA), and N-hydroxysuccinimide ester diazine. Salts are defined as chemicals that contain positively and negatively charged ions. Examples include calcium chloride and calcium sulfate. Calcium chloride (CaCl) is often added to the surface of implants to promote surface-initiated physical crosslinking. 2 ) or calcium sulfate (CaSO 4 ) to deposit other salts of divalent cations (e.g., Ca 2+ , Mg 2+ , Sr 2+ ), or other salts of polyvalent cations (e.g., Ti 4+ , Al 3+) may be used for surface initiated physical crosslinking. A hydrogel precursor solution is defined as a solution containing a crosslinkable polymer and an initiator. Charged polymers are defined as polymers that have a negative or positive charge, such as alginate or polyglutamic acid. · Covalently reactive macromonomers are defined as polymer molecules that contain chemically reactive groups. - Initiator is defined as a chemical that starts a chemical reaction. Examples include photoinitiators (e.g. lithium phenyl-2,4,6-trimethylbenzoylphosphinate) and chemical initiators (e.g. ammonium persulfate). A biological substance is defined as any molecule or organism that can interact with a living organism. Examples include proteins, peptides, cells, DNA, RNA, drugs, antibiotics, etc. A coating with one or more layers is defined as a coating with a single layer having a thickness in the range of 10-1000 μm or a coating with several layers, each having a thickness in the range of 10-1000 μm. · Tissue engineering is defined as the engineering or regeneration of hard or soft tissues such as bone, cartilage, tendons, ligaments, muscles, heart and heart valves.

[0047] The present invention provides a facile method for fabricating PCL-TCP / hydrogel composite scaffolds, hereafter referred to as hybrid tissue engineered constructs (HyTECs). After 3D printing, the surface of the PCL-TCP scaffolds was treated in three steps to increase hydrophilicity, improve hydrogel adhesion, and promote (induce) surface-initiated crosslinking. The scaffolds were then loaded with physically crosslinked hydrogels, which were then covalently crosslinked to form a stable interpenetrating network. The effects of surface treatments, process parameters, and freeze-drying on hydrogel loading and maintenance of the porous structure of the scaffolds after fabrication were investigated. The adhesion between PCL-TCP and hydrogels, and the release kinetics of BMP2 protein encapsulated in HyTECs were also investigated. Furthermore, the biocompatibility and osteoinductive potential of BMP2-loaded HyTECs were investigated. The disclosed method / strategy can be used to incorporate a wide range of hydrogels into porous polyester-based scaffolds and to coat non-porous polyester-based constructs. As an example, at the end of this specification, we demonstrate the effectiveness of the method / strategy to coat non-porous PCL-TCP rods by modifying the hydrogel and crosslinking mechanism for sustained release of BMP2 protein and cell encapsulation.

[0048] Materials and Methods

[0049] The following description of materials and methods is an exemplary embodiment.

[0050] material

[0051] Medical grade polycaprolactone (PCL, Mn = 80 kDa) was purchased from Sigma-Aldrich. β-TCP nanopowder (TCP) with an average particle size of 100 nm was obtained from Berkeley Advanced Materials Inc. N,N-dimethylformamide (DMF), sodium hydroxide (NaOH), and ethanol were purchased from Fisher Scientific Inc. N-(3-Dimethylaminopropyl)-N´-ethylcarbodiimide hydrochloride (EDC), N-hydroxysulfosuccinimide (NHS), 2-(N-morpholino)ethanesulfonic acid (MES), N-(3-aminopropyl)methacrylamide hydrochloride (APMA), ammonium persulfate (APS), N,N,N´,N´-tetramethylethylenediamine (TEMED), type A gelatin, heparin, ninhydrin, and TritonX-100 were purchased from Sigma-Aldrich. Ninhydrin reagent was prepared by dissolving 20 mg / mL ninhydrin in ethanol. Polyethylene glycol dimethacrylate (PEGDMA, Mn = 1000 g / mol) was obtained from Polyscience, Inc. Sodium alginate (alginic acid, 500GM) was purchased from Pfaltz & Bauer Inc. Human BMP2 protein was provided by Medtronic. Calcium colorimetric assay (MAK022), CCK-8 kit, and human BMP2 ELISA kit were purchased from Sigma-Aldrich. Quanti Chrom ALP assay kit was received from BioAssay Systems LLC. Quant-it PicoGreen assay kit was purchased from Fisher Scientific.

[0052] method

[0053] Synthesis and 3D Printing of PCL-TCP Filament

[0054] PCL-TCP filaments with a weight ratio of PCL to TCP of 80:20 were synthesized as described by Bruyas (Bruyas et al., Effect of Electron Beam Sterilization on Three-Dimensional-Printed Polycaprolactone / Beta-Tricalcium Phosphate Scaffolds for Bone Tissue Engineering, Tissue Eng Pt A (2018)). Briefly, 80 g of PCL and 20 g of TCP were dissolved in 800 mL and 400 mL of DMF, respectively, with stirring at 80 °C for 3 h. The PCL and TCP solutions were then mixed and stirred for 1 h before being precipitated in 4 L of water to produce the PCL-TCP composite. The PCL-TCP composite was washed with water and residual solvent was removed by air drying at ambient temperature for 24 h. Dried PCL-TCP composites were cut into pellets and extruded using a home-made screw extruder as described by Bruyas (Bruyas et al, Systematic characterization of 3D-printed PCL / beta-TCP scaffolds for biomedical devices and bone tissue engineering: Influence of composition and porosity, J Mater Res 33(14) (2018) 1948-1959). PCL-TCP scaffolds were 3D printed using a Lulzbot Mini (Aleph Objects Inc, USA) with a nozzle diameter of 500 μm. For surface characterization, non-porous disks with a diameter of 10 mm and a thickness of 600 μm were printed, and for all other tests, porous cylinders with a diameter of 10 mm and a height of 5 mm were printed. The 3D models were designed using SolidWorks (SolidWorks Corp.) and sliced ​​using Cura software.Post distances of 0.4, 0.53, 0.80, 1.00, 1.25, and 2.00 mm were used to print 0%, 30%, 50%, 60%, 70%, and 80% porous scaffolds. Printing temperature, layer thickness, and printing speed were set at 160 °C, 200 μm, and 5 mm / s, respectively, as described by Bruyas et al. (Effect of Electron Beam Sterilization on Three-Dimensional-Printed Polycaprolactone / Beta-Tricalcium Phosphate Scaffolds for Bone Tissue Engineering, Tissue Eng Pt A (2018)). To synthesize methacrylated gelatin (GelMA) macromonomer, gelatin was dissolved in deionized water (10% w / v) at 50 °C. Methacrylic anhydride was added to the gelatin solution in a molar ratio of 100:1 (methacrylic anhydride:gelatin) and reacted for 1 h at 50 °C with stirring. The mixture was then diluted 5-fold with deionized water and dialyzed against deionized water at 40 °C for 3 days using dialysis tubing with a molecular weight cutoff of 6–8 kDa (Spectrum Laboratories, Rancho Dominquez, CA, USA). The GelMA solution was then lyophilized and stored at -80 °C.

[0055] To synthesize methacrylated heparin (HepMA), 1 g of heparin was dissolved in 100 mL of MES buffer (100 mM). Subsequently, 5 mL of MES buffer containing 45 mg of EDC and 30 mg of NHS was added to the heparin solution to activate the carboxylic acid groups (Jeon et al, Affinity-based growth factor delivery using biodegradable, photocrosslinked heparin-alginate hydrogels, J Control Release 154(3) (2011) 258-66). After reacting for 1 h at room temperature, 25 mg of APMA in 1 mL of MES was added to the heparin solution and reacted for 2 h at room temperature. Next, the methacrylated heparin solution was dialyzed against deionized water at room temperature for 3 days using dialysis tubing with a molecular weight cutoff of 6–8 kDa (Spectrum Laboratories, Rancho Dominquez, CA, USA), and then lyophilized and stored at -80 °C.

[0056] Hydrogel loading on 3D printed porous scaffolds

[0057] The procedure of scaffold surface treatment and hydrogel formation is shown diagrammatically in FIG. 2A. The 3D printed scaffold was immersed in 5N NaOH solution and centrifuged at 1000 rpm for 1 min to ensure the penetration of the NaOH solution into the pores of the scaffold. Unless otherwise specified, the scaffold was incubated in the NaOH solution for 1 h and then washed with deionized water three times. Subsequently, the scaffold was incubated in MES buffer (100 mM) containing EDC (5 mg / mL) and NHS (5 mg / mL) for 30 min at room temperature to activate the carboxylic acid groups on the surface. Next, the scaffold was treated with APMA (0.25, 2.5, or 10 mg / mL) in MES (100 mM) buffer for 30 min at room temperature and then washed with deionized water. The surface of the APMA-treated scaffold was then treated with CaCl 2 The scaffold was wetted with a solution (20-200 mg / mL in deionized water) containing CaCl 2After incubation in the solution for 1 hour at room temperature, the scaffolds were centrifuged at 1000 rpm for 1 minute to remove residual solution. 2 The scaffolds with surface modification by the treatment were dried in vacuum for 3 hours and used for hydrogel loading.

[0058] The hydrogel precursor solution was prepared by dissolving PEGDMA (10–30% wt / vol), alginate (1.5% wt / vol), and 1 mg / mL APS in deionized water. The surface-treated scaffolds were immersed in the hydrogel precursor solution for 1 min at room temperature and then centrifuged at 1000 rpm for 1 min to remove the residual precursor solution. In this step, a hydrogel layer was formed on the scaffold surface due to the diffusion of calcium from the surface of the scaffold and the gelation of alginate. The hydrogel-coated scaffolds were then incubated in APS (9 mg / mL) and TEMED (6 mg / mL) in deionized water for 5 min to crosslink the PEGDMA macromonomers in the hydrogel layer and form an interpenetrating network. The scaffold / hydrogel complexes were washed with deionized water to remove any residual initiator or unreacted macromonomers.

[0059] Coating non-porous rods with bioabsorbable hydrogels

[0060] The procedure for coating non-porous PCL-TCP rods with bioabsorbable hydrogels is shown in Figure 18A. PCL-TCP filaments with a diameter of 0.9 mm were synthesized as described below, manually cut into 15 mm rods, and immersed in 5N NaOH solution for 6 h. The rods were then washed three times with deionized water and incubated in MES buffer (100 mM) containing EDC (5 mg / mL) and NHS (5 mg / mL) for 30 min at room temperature. The rods were then washed three times with deionized water and incubated in a 2% solution of gelatin methacrylate (GelMA) in MES buffer at 37 °C for 1 h. The rods were then washed three times with deionized water to remove unreacted GelMA and incubated in EDC / NHS (5 mg / mL) in MES buffer for 15 min at room temperature. The GelMA-coated rods were then washed three times with deionized water and dried under vacuum. The GelMA-coated rods were then immersed in CaSO in deionized water. 4 The hydrogel-loaded rods were immersed in the suspension at 60°C and sonicated for 30 seconds. The rods were then transferred to wells of a 24-well plate and dried under vacuum. The dried rods were immersed in wells of a 96-well plate containing GelMA (15%), alginate (1.25%), PEGDMA (2%), HepMA (1%), protein (BMP2; 200 μg / mL), and photoinitiator (0.3%) in deionized water for 2 minutes at 37°C. The hydrogel-loaded rods were removed from the solution and left in a dried well of another 96-well plate for 5 minutes. The hydrogel-loaded rods were then irradiated with visible light for 15 minutes to covalently crosslink GelMA, PEGDMA, and HepMA. The crosslinked hydrogel-loaded rods were stored at -80°C and lyophilized.

[0061] Freeze-drying and electron beam sterilization

[0062] For freeze-drying, HyTECs were immersed in liquid nitrogen for 30 min, then freeze-dried and stored at 4 °C. For rehydration, freeze-dried HyTECs were incubated in deionized water for 15 min and then further analyzed. For electron beam sterilization, HyTECs were exposed to electron beam irradiation at a standard single dose of 25 kGy according to standard ISO11137-2:2006. Electron beam sterilization of PCL-TCP rods loaded with BMP2-containing hydrogels was performed with one dose (25 kGy) or two doses (12.5 kGy) to investigate the effect of fractionated electron beam irradiation on the activity of the loaded protein.

[0063] Surface and mechanical properties evaluation

[0064] The density of grafted APMA on the PCL-TCP scaffold was quantified by measuring the unreacted APMA concentration in the solution after reaction using a ninhydrin assay. Briefly, the APMA solution after reaction with the PCL-TCP scaffold was diluted 10-fold with MES buffer. To 200 μL of the diluted APMA solution, 40 μL of ninhydrin reagent was added. After mixing, the solution was heated to 90 °C for 8 min in a capped tube and the absorbance was read at 570 nm using a SpectraMax M2 plate reader (Molecular Devices LLC). The unreacted APMA concentration in the solution was calculated using a calibration curve constructed against the absorbance of APMA solutions of known concentrations. To evaluate the surface hydrophilicity of the PCL-TCP constructs, a 4 μL water droplet was deposited on the disk scaffold and the contact angle was measured using a Rame-Hart 290 goniometer (Rame-Hart instrument co., USA) and analyzed using image processing.

[0065] The apparent Young's modulus and yield stress of the scaffolds were tested using an Instron 5944 uniaxial testing system (Instron Corporation, Norwood, Massachusetts, USA) with a 2 kN load cell, 1 N preload, and a displacement rate of 1% strain / sec. The initial slope of the stress-strain curve was taken as the Young's modulus. The yield stress was defined as the stress at which a line starting at 1% strain offset with a slope equal to the Young's modulus intersects the stress versus strain curve.

[0066] To measure the release of calcium ions from the surface of the scaffold, CaCl 2 Scaffolds treated with CaSO 4 The Ca release medium (non-porous rods treated with Ca) was incubated in 1 mL of deionized water at room temperature for 1 h. 2+ Ion concentrations were measured at 575 nm using a calcium colorimetric assay (MAK022; Sigma-Aldrich, USA) on a SpectraMax M2 plate reader (Molecular Devices, Inc.).

[0067] To measure the thickness of the hydrogel layer, the scaffolds were imaged before and after hydrogel loading using a Dino-Lite digital microscope camera. The images were then analyzed using ImageJ to quantify the average hydrogel thickness. The average thickness of the gel in the square pores was defined as half the difference in size of the voids (black in the images) before and after hydrogel loading. The percentage of pores filled with hydrogel was defined as the ratio of the number of square pores completely filled with hydrogel to the total number of square pores.

[0068] The hydrogel loading rate (%) was calculated from the scaffold weight (Wb) before hydrogel loading and the scaffold weight (Wa) after hydrogel loading using the following formula. Hydrogel loading rate (%) = 100 × (Wa-Wb) / Wb

[0069] For scanning electron microscopy (SEM) imaging, HyTEC samples were immersed in liquid nitrogen and freeze-dried. The freeze-dried samples were immersed in liquid nitrogen and cut with a scalpel. The hydrogel samples were subsequently coated with gold for 180 s using SPI sputtering (SPI Supplier Division of Structure Prob, Inc., West Chester, PA, USA) and imaged using a field emission scanning electron microscope (Zeiss Sigma, White Plains, NY, USA) at an accelerating voltage of 5 keV.

[0070] Hydrogel interface stiffness

[0071] A customized 3D printed PCL-TCP device was designed and used to evaluate the adhesion of hydrogel to PCL-TCP scaffolds (see Fig. 16A, Fig. 16B). The device consists of two concentric cylinders (inner and outer cylinders) separated by a gap of 400 μm and connected via two bridges. Hydrogel was incorporated into the gap between the two concentric cylinders. The hydrogel-incorporated device was then placed on an Instron 5944 uniaxial testing system (Instron, Norwood, MA, USA) equipped with a 100 N load cell, and the bridge connecting the inner and outer cylinders was cut (see Fig. 16C). The force required to push the inner cylinder out of the device (interface stiffness) was measured by a push-out test using a preload of 0.1 N and a displacement rate of 0.1 mm / s.

[0072] Protein release

[0073] To measure the release kinetics from porous HyTEC, BMP2 protein was added to PEGDMA (10–30% wt / vol), alginate (1.5% wt / vol), and APS (1 mg / mL) precursor solutions prior to hydrogel loading on 80% porous PCL-TCP scaffolds. Increasing the PEGDMA concentration from 10% to 20% or 30% changed the average hydrogel loading (relative to scaffold weight) from 151% to 169% or 144% (see Figure 16F). To load 1.5 μg of BMP2 on all scaffolds, 9.0, 8.0, and 9.5 μg / mL of BMP2 were added to the 10%, 20%, and 30% PEGDMA precursor solutions, respectively. BMP2-loaded HyTECs were incubated in 1 mL of PBS at 37 °C for 35 days. At each time point, the amount of BMP2 in the release medium was measured using ELISA, and the release medium was replaced with 1 mL of fresh PBS.

[0074] To measure the kinetics of BMP2 release from nonporous HyTECs, rod-shaped HyTECs encapsulating 2 μg of BMP2 were freeze-dried and incubated in 1 mL of PBS at 37 °C for 28 days. At each time point, the amount of BMP2 in the release medium was measured using ELISA, and the release medium was replaced with fresh PBS.

[0075] cell culture

[0076] Human mesenchymal stem cells (hMSCs) and pluripotent mouse C3H10T1 / 2 fibroblasts (ATCC, USA) were cultured in DMEM medium supplemented with 10% fetal bovine serum (FBS; Life Technologies, USA), 1% penicillin, and streptomycin (hereafter referred to as culture medium; Life Technologies, USA) at 37 °C and 5% CO. 2 The cells were cultured in a humidified incubator at 4 °C for 1 h. After reaching 70% confluency, hMSC or C3H10 cells were enzymatically suspended with trypsin-EDTA and used for in vitro studies. All cells were passaged less than six times before in vitro studies.

[0077] Biocompatibility of Porous HyTEC

[0078] For in vitro cell testing, after surface modification with APMA and CaCl 2 The untreated 80% porous PCL-TCP scaffolds were sterilized with 70% ethanol solution for 20 min.

[0079] CaCl 2 The solutions and hydrogel precursor solutions were sterilized by filtration using a 0.22 μm Millex syringe filter.

[0080] Scaffolds with or without hydrogel loading were cultured in 1 mL of culture medium at 37° C. Medium without scaffolds incubated at 37° C. was used as the control group. The incubated cell culture medium (hereafter referred to as conditioned medium) on day 1 (for viability and proliferation tests) and day 4 (for proliferation tests) was used for viability and proliferation tests.

[0081] For cell viability tests, scaffolds or scaffold / hydrogels were incubated in culture medium while hMSCs and C3H10 cells were seeded at 5000 cells / well in 96-well plates and incubated at 37°C, 5% CO. 2 The cells were incubated at 4°C for 24 h. The culture medium was then replaced with 100 μL of conditioned medium, and the cells were incubated for another 24 h. To measure cell viability, 10 μL of CCK-8 solution (CCK-8 kit; Sigma-Aldrich) was added to each well, and the absorbance was read at 450 nm on a plate reader after 3 h of incubation. The viability of cells in the experimental groups (scaffold, scaffold + hydrogel) was divided by the viability of cells in the control group (no scaffold) to calculate the normalized viability.

[0082] For cell proliferation studies, scaffolds or scaffold / hydrogels were incubated in culture medium while hMSCs and C3H10 cells were seeded at 10,000 cells / well in 24-well plates and incubated at 37°C, 5% CO. 2The cells were cultured at 4°C for 24 h. The culture medium was then replaced with 600 μL of conditioned medium, and the cells were incubated for an additional 7 days. Three days after addition of conditioned medium (four days after cell seeding), the medium was replaced with fresh conditioned medium. On days 0, 3, and 7, the cells were washed with PBS, enzymatically suspended using 250 μL of 0.25% trypsin-EDTA solution (Life Technologies, USA), and counted using a Z2 particle counter (Beckman Coulter, USA).

[0083] Osteoinductive potential of BMP2-loaded porous HyTEC

[0084] HyTECs without BMP2 (scaffold + hydrogel) or with 1.5 μg of BMP2 (scaffold + hydrogel / BMP2) were incubated in 1 mL of medium at 37 °C. The conditioned medium was used for cell differentiation tests and replaced with fresh medium every 3 days. Medium without scaffold and BMP2 (control), medium supplemented with 1.5 μg / mL of BMP2 (BMP2 in medium (3d)), and medium supplemented with 214 ng / mL of BMP2 for 21 days (BMP2 in medium (21d)) incubated at 37 °C were used as control groups. At the same time that the experimental and control groups were incubated in culture medium, hMSCs and C3H10 cells were seeded at 10,000 cells / well in 24-well plates and incubated at 37 °C, 5% CO. 2The cells were incubated at 4°C for 24 h. The culture medium was then replaced with 600 μL of conditioned medium, and the cells were incubated for another 21 days, with the medium being replaced with fresh conditioned medium every 3 days. For BMP2 in medium (3d) group, the medium was changed to medium (BMP2-free) after 3 days. At each time point (days 0, 7, 14, and 21), the cells were washed with PBS and lysed with 1% TritonX-100 in PBS using a cell scraper, followed by shaking at room temperature for 20 min. The lysate was centrifuged at 2000 × g for 15 min at 4 °C, and the supernatant was collected. ALP activity in the supernatant was measured at 405 nm on a plate reader using the QuantiChrom ALPA Assay Kit (BioAssay Systems, Hayward, CA, USA) according to the manufacturer's instructions. The double-stranded DNA content of the lysates was measured using a PicoGreen assay kit (Quant-it; Thermo Fisher Scientific). Normalized ALP activity was calculated by dividing the ALP activity by the DNA content.

[0085] Cell-loaded HyTEC

[0086] PCL-TCP filaments with a diameter of 0.9 mm were synthesized and coated with GelMA as described in the previous section. The GelMA-coated rods were then heated with CaSO in deionized water at 60 °C. 4The rods were then immersed in a hydrogel precursor solution containing GelMA (10%), alginate (1.25%), PEGDMA (2%), and photoinitiator (0.3%) in calcium-free medium at a density of 2 million cells / mL. The sterile rods were then immersed in the cell precursor solution for 2 min at 37 °C. The rods carrying the cell-laden hydrogel were removed from the solution and left in a sterile dry well of a 96-well plate for 5 min, followed by irradiation with visible light for 15 min. The cell-laden hydrogel-coated rods were then transferred to a well of a 24-well plate and sterilized in culture medium at 37 °C and 5% CO. 2 and incubated at 4°C for 1 hour.

[0087] For live / dead cell imaging, cell-loaded HyTECs were stained with calcein AM (2 μM) and ethidium homodimer 1 (4 μM) according to the manufacturer's instructions and imaged using a Zeiss Axio Observer Z1 fluorescent microscope. Live / dead cell images were divided into small squares and the number of live and dead cells was manually counted to calculate cell viability. To quantify the DNA content of HyTECs, at each time point, samples were transferred to new wells and incubated for 1 h at 37 °C in 500 μL of DMEM medium supplemented with collagenase (1 mg / mL). Subsequently, 250 μL of Triton solution (3%) in PBS was added to each well and attached cells were scraped off from the surface using a CytoOne cell scraper (USA Scientific Inc, Ocala, FL, USA). The cell suspension was then transferred to a microcentrifuge tube and sonicated. The cell lysates were then centrifuged at 2000 × g for 15 min at 4 °C and the supernatants were collected. The double-stranded DNA content in the supernatants was measured using the Quant-iT PicoGreen DNA assay according to the manufacturer's instructions.

[0088] statistical analysis

[0089] All experiments were performed in triplicate. Statistically significant differences between groups were tested using a two-way repeated measures analysis of variance (ANOVA) followed by a two-tailed Student's t-test. A p-value less than 0.05 (p<0.05) was considered statistically significant.

[0090] result

[0091] The density of grafted APMA on PCL-TCP constructs versus APMA concentration in the reaction solution is shown in Figure 11A. When the APMA concentration increased from 0.25 to 2.5 mg / mL, the density of grafted APMA increased from 7.4 to 13.6 (μg / mg scaffold). When the APMA concentration increased from 2.5 to 10 mg / mL, no significant change in the density of grafted APMA was observed. The effect of surface modification with APMA on the contact angle of PCL-TCP constructs is shown in Figure 11B. The contact angle of the PCL-TCP construct without NaOH treatment (untreated) was 105.6°, indicating a hydrophobic surface. After NaOH treatment, the contact angle of the constructs decreased to 73.1°. This indicates that the PCL-TCP surface became relatively hydrophilic after NaOH treatment due to the detachment of hydroxyl and carboxyl groups from the PCL ester groups. The contact angle of the PCL-TCP constructs, i.e., the hydrophilicity of the PCL-TCP constructs, was not significantly different when the APMA concentration was 0.25 (A-0.25), 2.5 (A-2.5), or 10 (A-10) mg / mL. The 3D printed porous scaffolds were treated with 2.5 mg / mL APMA solution unless otherwise specified. PCL-TCP scaffolds with porosity ranging from 0% to 80% (Figure 11C) were fabricated to investigate the effect of surface treatment on the mechanical properties of the scaffolds. The Young's modulus and yield stress of the PCL-TCP scaffolds without surface treatment (Untreated; B) and the PCL-TCP scaffolds with APMA grafting after NaOH treatment (Treated / A-2.5; R) are shown in Figure 11D and Figure 11E, respectively. As the porosity increased from 0% to 30%, 60%, or 80%, the Young's modulus of the untreated scaffold decreased from 134.7 to 83.0, 42.5, or 21.2 MPa. With surface modification, the Young's modulus of the porous scaffold with 0% porosity decreased from 134.7 to 123.7 (MPa). No significant differences in the Young's modulus of the untreated and treated scaffolds were observed at 30%, 60%, or 80% porosity. As the scaffold porosity increased from 0% to 80%, the yield stress of the untreated scaffold decreased from 9.8 to 1.3 MPa.The yield stress of the scaffolds did not change significantly with surface treatment at any porosity.

[0092] Figure 12A shows the effect of NaOH treatment time on hydrogel coating on porous PCL-TCP scaffolds with 80% porosity. The thickness of the hydrogel layer and the percentage of filled pores are shown in Figure 12A and Figure 12D, respectively. The average hydrogel thickness increased from 94 to 248, 418, or 602 (μm) when the NaOH treatment time increased from 0 to 60, 120, or 180 min. No significant difference was observed in the percentage of filled pores between untreated scaffolds (0 min) and scaffolds treated with NaOH solution for 60 min. The percentage of filled pores increased significantly with increasing NaOH treatment time from 60 to 120 to 180 min (see Figure 12D). The effect of NaOH treatment time on the release of calcium ions from the surface of the scaffolds is shown in Figure 12B. With increasing NaOH treatment time from 0 to 180 min, the calcium concentration in the release medium increased monotonically from 0.24 to 0.98 (mg / mL). Therefore, the increase in hydrogel thickness and the percentage of filled pores with increasing NaOH treatment time may be due, in part, to the release of a large amount of calcium from the scaffolds.

[0093] CaCl in the incubation solution for hydrogel coating of porous PCL-TCP scaffolds with 80% porosity 2 The effect of concentration is shown in Figure 13A. The thickness of the hydrogel layer and the percentage of filled pores are shown in Figures 13C-D, respectively. 2 When the concentration was in the range of 20-100 mg / mL, the average thickness of the hydrogel was less than 259 μm and the percentage of filled pores was 6.7%, showing no statistically significant change. 2As the concentration increased from 100 to 200 mg / mL, the average thickness of the hydrogel increased from 259 μm to 519 μm, and the percentage of filled pores increased from 6.7% to 31.8%. The effect of CaCl in the incubation solution on the release of calcium ions from the surface of the scaffolds. 2 The effect of concentration is shown in Figure 13B. 2 As the concentration increased from 20 to 200 mg / mL, the released Ca in the release medium 2+ The concentration increased three-fold. 2 The dramatic change in hydrogel thickness and percentage of filled pores by increasing the concentration from 100 to 200 mg / mL (see Figure 13B-C) indicates that Ca 2+ This is believed to be due to a significant increase in the amount of released (see FIG. 13D).

[0094] 3D printed porous PCL-TCP scaffolds with 50%, 60%, or 70% porosity before and after hydrogel coating are shown in Figure 14A. The effect of scaffold porosity on hydrogel thickness and percentage of filled pores is shown in Figure 15B-C. As scaffold porosity increased from 50% to 80%, the average hydrogel thickness increased from 128 to 248 μm and the percentage of filled pores decreased from 23.9% to 8.4%. CaCl 2 (100 mg / mL) scaffold with incubation in solution (CaCl 2 Treatment; R) and CaCl 2 Scaffolds without incubation in solution (CaCl 2 The effect of scaffold porosity on calcium ion release from the surface of untreated (B) is shown in Figure 14D. 2 Without incubation, the released Ca in the release medium 2+ The average concentration of Ca was less than 0.07 mg / mL, and no significant change was observed when the porosity of the scaffold was changed from 50% to 80%. When the porosity was increased from 50% to 80%, the Ca concentration from the scaffold was increased. 2+The release of Ca decreased from 0.62 to 0.50 mg / mL, but the Ca release per unit weight of the scaffold was 2+ The amount of released increased from 3.03 to 5.02 (mg / mg scaffold) (Figure 14D-E). In addition, the hydrogel loading (relative to the scaffold weight) increased from 84% to 165% when the porosity increased from 50% to 80% (Figure 14F).

[0095] The effect of freeze-drying on the mechanical properties of PCL-TCP scaffolds and the properties of the hydrogel layer on the scaffolds are shown in Figures 15A-G. Freeze-drying increased the Young's modulus, stress, and yield of PCL-TCP scaffolds at all porosities (Figures 15A-B). The hydrogel layer remained intact after freeze-drying and rehydration (Figure 15C). In addition, freeze-drying and rehydration had no significant effect on the thickness and percentage of filled pores of the hydrogel, regardless of the NaOH treatment time (Figures 15D-E). CaCl in the incubation medium 2 At concentrations below 100 mg / mL, the thickness of the hydrogel layer and the percentage of filled pores did not change significantly upon lyophilization / rehydration. 2 At a concentration of 200 mg / mL, the hydrogel layer thickness and the percentage of filled pores of the freeze-dried / rehydrated scaffolds were significantly higher than those of the non-freeze-dried scaffolds.

[0096] The structure of the 3D printed PCL-TCP device used to measure the adhesion of hydrogel to the scaffold is shown in Figure 16A-B. The hydrogel was incorporated into the gap between two concentric cylinders. The hydrogel-incorporated device was then placed on an Instron instrument, the two bridges connecting the inner and outer cylinders were cut, and the interfacial stiffness was measured by a push-out test (Figure 16C). When the scaffold surface was treated with APMA, the interfacial stiffness of the hydrogel increased two-fold from 2.84 to 5.60 (N / mm) (Figure 16D). In addition, the interfacial stiffness of the hydrogel increased significantly from 2.72 to 5.60 or 10.24 (N / mm) when the PEGDMA concentration in the hydrogel solution was increased from 10% to 20% or 30% (Figure 16E). In contrast to the interfacial stiffness, the hydrogel loading on the scaffold was not significantly changed by increasing the PEGDMA concentration (Figure 16F). The effect of PEGDMA concentration on the release kinetics of enzymatically active BMP2 protein from porous HyTEC scaffolds is shown in Figure 16G. After an initial burst release, BMP2 was released steadily from hydrogel-loaded scaffolds over 7 days, followed by a lower rate from days 7 to 35. The enzymatically active BMP2 protein released from HyTEC after 35 days increased from 28% to 44% or 61% when the PEGDMA concentration was decreased from 30% to 20% or 10%, respectively. The effect of freeze-drying and e-beam sterilization on the release kinetics of BMP2 from PCL-TCP / hydrogel scaffolds is shown in Figure 16H. The BMP2 release from freeze-dried scaffolds after 35 days was 26% lower than that of untreated scaffolds. In addition, the BMP2 release from freeze-dried and e-beam irradiated scaffolds was 22% lower than that from scaffolds that were not freeze-dried and e-beam irradiated after 35 days.

[0097] The normalized viability and proliferation of hMSCs and C3H10 cells cultured in DMEM medium pretreated with PCL-TCP / hydrogel or pristine PCL-TCP scaffolds are shown in Figure 17A-C. The viability of hMSCs or C3H10 cells cultured in scaffold / hydrogel pretreated medium for 24 h was not significantly different from that of cells cultured in scaffold pretreated medium (Figure 17A). The cell number of hMSCs in the scaffold / hydrogel group was 2.6 × 10 on day 0. 4 12.5 × 10 on the 3rd and 7th days, respectively 4 pcs and 24.0×10 4 The proliferation of hMSCs in the scaffold / hydrogel or scaffold group was not statistically different from the control (no scaffold) group at any time point ( FIG. 17B ). The number of C3H10 cells in the scaffold / hydrogel group was 2.2×10 on day 0. 4 49.0×10 on the 7th day 4 The number of C3H10 cells increased to 100 cells / ml (FIG. 17C). The C3H10 proliferation in the scaffold / hydrogel or scaffold group was not statistically significantly different from the control (no scaffold) group at any time point. Therefore, the process of surface treatment and hydrogel formation did not adversely affect cell viability and proliferation.

[0098] The ALP activity of hMSCs and C3H10 cells cultured in DMEM medium (control), DMEM medium pretreated with scaffold / hydrogel without BMP2 (scaffold+hydrogel), DMEM medium pretreated with scaffold+hydrogel containing BMP2 (scaffold+hydrogel / BMP2), DMEM medium supplemented with 1.5 μg / mL of BMP2 for 3 days (no scaffold / BMP2 (3d)), and DMEM medium supplemented with 214 ng / mL of BMP2 for 21 days (no scaffold / BMP2 (21d)) is shown in Figure 17D-E. The ALP activity of hMSCs in the control and scaffold+hydrogel groups did not increase significantly over 21 days. The ALP activity of hMSCs in the scaffold+hydrogel / BMP2, no scaffold / BMP2 (3d), and no scaffold / BMP2 (21d) groups increased significantly after 14 days of culture, and then did not increase significantly from days 14 to 21. On days 14 and 21, the ALP activity of hMSCs in the scaffold+hydrogel / BMP2 group was significantly higher than that in the no scaffold / BMP23d group and significantly lower than that in the no scaffold / BMP221d group. The ALP activity of the control and scaffold+gel groups did not change significantly over 21 days of culture. The ALP activity of C3H10 in the no scaffold / BMP2 (3d) group increased from day 0 to day 7, and then decreased. The ALP activity of C3H10 in the scaffold + hydrogel / BMP2 group and the no scaffold / BMP2 (3d) group increased significantly from day 0 to day 14, and no significant change was observed from day 14 to day 21. On days 14 and 21, the ALP activity of C3H10 cells in the scaffold + hydrogel / BMP2 group was significantly higher than that in the no scaffold / BMP2 (3d) group and significantly lower than that in the no scaffold / BMP2 (21d) group.The ALP activity in the no scaffold / BMP2(3d) group was significantly higher than that in the scaffold+hydrogel / BMP2 or no scaffold / BMP2(21d) groups at day 7, but the peak ALP activity in the no scaffold / BMP2(3d) group (6.1 IU / mg DNA) over 21 days was significantly lower than the peak ALP activity in the scaffold+hydrogel / BMP2 (10.3 IU / mg DNA) or no scaffold / BMP2(21d) groups (25.1 IU / mg DNA).

[0099] To fabricate non-porous HyTEC constructs, NaOH treatment was followed by freeze / thawing, binding of GelMA to the surface, and addition of CaSO. 4 The deposition increased hydrophilicity, improved hydrogel adhesion, and promoted (induced) surface-initiated crosslinking. SEM images of the surface of the hydrogel supported on the non-porous PCL-TCP rods are shown in Figure 18B-C. The surface of the hydrogel was porous, with an average pore size of 8 μm. Ca from the surface of the PCL-TCP rods was 2+ CaSO in the suspension used for scaffold treatment at 60 °C for the release of 4 The effect of concentration was examined by immersing rods in 10 mg / mL or 20 mg / mL CaSO without freeze / thawing. 4 No release of calcium was detected when treated in suspension. CaSO in suspension, with or without freezing / thawing 4 When the concentration was increased from 50 mg / mL to 200 mg / mL, the amount of calcium released from the rod surface significantly increased. 4 The amount of calcium released from the surface of the rods increased significantly when the concentration was 20, 50, or 100 mg / mL. 4 When the effect of temperature on the suspension (50 mg / mL) was examined, calcium release was significantly increased by CaSO 4The calcium deposition on the PCL-TCP rods was significantly increased by increasing the temperature of the suspension from 25 to 60 °C, and did not change with further increase from 60 to 70 °C. For the BMP2-containing hydrogel coating, calcium deposition on the PCL-TCP rods was significantly increased by 50 mg / mL CaSO 4 The results were analyzed in suspension at 60°C. The average loading rates of BMP2-containing and BMP2-free hydrogels on treated calcium-deposited PCL-TCP rods were 111% and 108%, respectively, with no significant difference between the two groups (Figure 18D). The release kinetics of BMP2 from freeze-dried BMP2-loaded nonporous HyTEC over 28 days is shown in Figure 18E. The amount of BMP2 released from HyTEC increased continuously over 14 days, with no significant change observed from day 14 to day 28. The total amount of BMP2 released from freeze-dried HyTEC after 28 days was 80.0%. The release kinetics of BMP2 from nonporous HyTEC after freeze-drying and sterilization using single or fractionated e-beam irradiation is shown in Figure 18E (B and R). The total amount of BMP2 released from freeze-dried HyTEC after sterilization using single or fractionated electron beam irradiation was 27.8% and 54.5%, respectively, after 28 days. Figures 18F-G show live / dead cell images of cell-loaded nonporous HyTEC and hMSC in HyTEC, respectively. The viability of hMSC cells in HyTEC ranged from 92% to 96% over 7 days of culture (Figure 18H), whereas the DNA content of cellular HyTEC increased 2.4-fold over 14 days of culture (Figure 18I).

[0100] Consideration

[0101] Complications associated with autografts, allografts, and DBM for the treatment of large bone defects underscore the importance of developing synthetic bone grafts. Many studies have demonstrated that PCL-TCP scaffolds are biocompatible, bioresorbable, mechanically stable, and osteoconductive. In addition, because PCL has a low melting point and is easy to process, PCL-TCP scaffolds can be fabricated by fused deposition modeling (FDM)-based 3D printing. However, 3D-printed PCL-TCP scaffolds lack bone growth-promoting proteins, limiting their application in the treatment of large bone defects. In this study, we developed a post-treatment method to fabricate 3D-printed PCL-TCP scaffolds that are interconnected, porous, and coated with a thick hydrogel layer containing proteins. After 3D printing, the surface of the scaffold can be treated in three successive steps to increase hydrophilicity, improve hydrogel adhesion, and promote (induce) surface-initiated crosslinking. The polyester surface was made hydrophilic by cleaving the ester bonds to the carboxyl and hydroxyl groups by treatment with NaOH. In addition, reactive double bonds were introduced to the surface of the scaffold by grafting APMA to the carboxyl groups using carbodiimide chemistry. To control the thickness of the hydrogel layer, the APMA-modified scaffold was immersed in CaCl 2 The surface-initiated crosslinking was promoted (induced) by treatment with CaCl 2 When the scaffold surface-treated with was immersed in a hydrogel precursor solution, CaCl was deposited on the surface of the scaffold. 2diffused from the surface of the scaffold into the solution and crosslinked the alginate near the surface, forming a hydrogel layer on the surface of the scaffold. The PEGDMA macromonomers in the physically crosslinked hydrogel were covalently crosslinked in the next step to form a rigid interpenetrating network. The hydrogel network was attached to the scaffold surface by the reaction of the double bonds of the PEGDMA macromonomers with the double bonds of the APMA grafted to the scaffold surface. The covalent bonding of the reactive functional groups of PEGDMA with the reactive functional groups on the scaffold surface during the crosslinking reaction improved the adhesion of the hydrogel to the surface of the PCL-TCP scaffold. The bonding of the functional groups on the surface to the polymer network improves the adhesion of the polymer network to a rigid surface. For example, the adhesion between PEGDA / alginate IPN hydrogels and glass, ceramic, titanium, or aluminum was significantly improved by modifying the surface with reactive 3-(trimethoxysilyl)propyl methacrylate (TMSPMA) and covalently anchoring the hydrogel to the surface.

[0102] The thickness of the hydrogel layer was directly correlated with the total amount of calcium ions released from the surface of the scaffold and was correlated with the NaOH treatment time and the CaCl in the treatment solution. 2 The increase in the total amount of calcium ions released from the scaffold with increasing NaOH treatment time is due to the improved hydrophilicity and roughness of the scaffold surface, i.e., the CaCl 2 This is due to the increased absorbency of the solution. Similarly, the CaCl 2 The increase in the total amount of calcium ions released from the scaffold with increasing concentration is due to increased deposition of calcium on the surface of the scaffold.

[0103] Despite the increase in hydrogel loading, the hydrogel thickness did not change dramatically with increasing scaffold porosity. Thus, the pore size of porous HyTEC can be adjusted by changing the porosity of the scaffold. The pore size of the interconnecting pores in bone tissue engineering scaffolds should be at least 100 μm for cell infiltration, bone ingrowth, and angiogenesis. Also, cell migration and bone ingrowth are optimal when the pore size is larger than 100 μm. For example, osteoblast adhesion and proliferation on collagen-glycosaminoglycan (CG) scaffolds with pore sizes greater than 300 μm were higher than those with pore sizes less than 200 μm

[24] . When porous poly(ether ester) block copolymer scaffolds were implanted into the dorsal skinfold chambers of BALB / c mice, vascular ingrowth was faster in scaffolds with large pores (250-300 μm) compared to scaffolds with medium pores (75-212 μm) or small pores (20-75 μm). In the present invention, the scaffold pore size after hydrogel loading ranged from 140 μm to 300 μm, 480 μm, or 1100 μm when the porosity of the pristine scaffold increased from 50% to 60%, 70%, or 80%. Although the pore size of all hydrogel-loaded scaffolds was larger than 100 μm, according to the aforementioned published reports, scaffolds with porosity of 60% or more and pore size of 300 μm or more after hydrogel loading would be optimal for future in vitro experiments.

[0104] The results presented here showed that BMP2 was released from the porous PCL-TCP scaffold loaded with alginate / PEG-based hydrogel over a period of 35 days. BMP2-loaded alginate / PEG-based hydrogel with sustained release of BMP2 has been shown to promote ectopic bone nodule formation in mice in vitro. Thus, the incorporation of BMP2-loaded alginate / PEG-based hydrogels imparts osteoinductivity to the osteoconductive 3D-printed porous PCL-TCP scaffolds, which can enhance and promote bone formation. At higher PEGDMA concentrations (Figure 16G), the lower amount of released protein was attributed to the increased crosslink density and decreased mesh size of the polymer network.

[0105] Lyophilization facilitates the storage / transportation of bioactive products and extends their shelf life. The results presented here show that lyophilization does not adversely affect the mechanical properties of the scaffold and the properties of the hydrogel layer, and reduces the release of bioactive BMP2. Since PCL-TCP scaffolds are heat sensitive, e-beam irradiation is considered a reliable method for terminal sterilization of heat sensitive materials. It was also shown that e-beam sterilization does not adversely affect the mechanical properties and degradation rate of PCL-TCP scaffolds. The present results reveal that e-beam sterilization reduces the release of bioactive BMP2 from freeze-dried BMP2-loaded porous HyTEC.

[0106] Furthermore, the results presented here showed that osteogenic differentiation of hMSCs and C3H10 cells was higher when cells were exposed to BMP2 released from porous HyTEC compared to cells exposed to BMP2 dissolved in culture medium at the same amount as when BMP2 was loaded onto hydrogels for 3 days. The higher osteogenic differentiation of cells exposed to BMP2-releasing scaffolds compared to cells exposed to BMP2 for 3 days was attributed to the time-dependent osteoinductivity of BMP2. The osteoinductivity of BMP2 protein is dose- and time-dependent. For example, the ALP activity of hMSCs exposed to slow BMP2-releasing electrospun PCL / PEG mats was significantly higher than that of hMSCs exposed to fast BMP2-releasing mats.

[0107] The method of the present invention can be used to load a wide range of hydrogels onto porous or non-porous polyester-based constructs. Thus, in addition to porous scaffolds, we investigated the effectiveness of the method to fabricate non-porous HyTEC using bioabsorbable hydrogels for sustained release of BMP2 protein. To improve the integration of the hydrogel with the non-porous rods, NaOH treatment was followed by freeze-thawing to increase the surface roughness. The results showed that freeze-thawing increased calcium deposition on the surface of the rods (not shown). GelMA was used as the macromonomer and PEGDMA was used as the crosslinker. HepMA was used to extend the release period of BMP2 due to the high affinity of heparin for BMP2. The addition of HepMA to alginate-based hydrogels was found to extend the release period of BMP2, thereby improving subcutaneous bone formation in mice. In addition, optical initiation was used for covalent crosslinking of the hydrogel on the non-porous rods instead of chemical initiation used for crosslinking of the hydrogel on the porous scaffolds. Chemically initiated crosslinking was used for porous scaffolds because it may limit light penetration to the central part of the scaffold. The release of BMP2 (as a model protein) from non-porous HyTEC demonstrated that the method can be used for sustained protein delivery with non-porous implants. Only 28% of the total amount of enzymatically active BMP2 was released from the hydrogel-loaded rods after freeze-drying and a single e-beam sterilization. This may be due to the denaturation of BMP2 or the GelMA-based network crosslinking with BMP2 under very strong radiation. However, splitting the high-intensity e-beam irradiation (25 kGy) into two low-intensity irradiations (12.5 kGy) nearly doubled the amount of active BMP2 released. The inventors also demonstrated that the disclosed method can be used to load hydrogels containing live cells onto scaffolds. The biodegradable hydrogel layer is thick enough to accommodate cells, and the cells are viable and able to proliferate. The methods of the present disclosure are particularly useful for loading cells onto implants.

Claims

1. A method for fabricating tissue engineering structures, (a) A step of treating the surface of the scaffold in order to increase the surface area of ​​the scaffold, (b) The step of depositing salt on the surface of the scaffold, (c) The scaffold obtained in step (b) is treated with a hydrogel precursor solution containing a biological substance, a charged polymer, a covalently reactive macromonomer, and an initiator to form a physically crosslinked hydrogel network on the surface of the scaffold. (d) A step of chemically crosslinking the covalently reactive macromonomers to form a tissue engineering construct, (e) A method comprising the step of freeze-drying the tissue engineering construct.

2. The method according to Claim 1, The scaffold is porous, in this method.

3. The method according to Claim 1, The method wherein the scaffold comprises polycaprolactone-β-tricalcium phosphate (PCL-TCP).

4. The method according to Claim 1, The method comprising step (a) treating the scaffold with a base, an acid, freeze / thaw, plasma, or any combination thereof.

5. The method according to Claim 1, Step (a) is a method for increasing the hydrophilicity or roughness of the surface of the scaffold.

6. The method according to Claim 1, The method wherein the charged polymer comprises an alginate.

7. The method according to Claim 1, The method comprises, for example, gelatin methacrylate (GelMA) and polyethylene glycol dimethacrylate (PEGDMA) as the covalently reactive macromonomer.

8. The method according to Claim 1, The method wherein the salt comprises Ca²⁺, Mg²⁺, Sr²⁺, Zn²⁺, Ti⁴⁺, or Al³⁺.

9. The method according to Claim 1, The method wherein the salt contains Ca²⁺.

10. The method according to Claim 1, The method wherein the salt comprises CaSO₄ or CaCl₂.

11. A tissue engineering structure, (a) A scaffold having a treated surface, (b) A hydrophilic hydrogel network physically crosslinked on the treated surface, (c) A biological substance contained within the physically crosslinked hydrophilic hydrogel network, (d) comprising a chemically crosslinked hydrogel network, Freeze-dried tissue engineering constructs.

12. The structure according to claim 11, The aforementioned scaffold is a porous, tissue-engineered structure.

13. A structure according to claim 11, The treated surface is a tissue-engineered construct obtained as a result of treatment with a base, acid, freeze / thaw, plasma, or any combination thereof.

14. A structure according to claim 11, A tissue engineering construct in which the physical crosslinking of the aforementioned component (b) is carried out via a charged polymer, salt ions, or salt.

15. The structure according to claim 14, The charged polymer is a tissue engineering construct containing alginate.

16. The structure according to claim 11, The chemically crosslinked hydrogel network is a tissue engineering construct located within the physically crosslinked hydrogel network.

17. A structure according to claim 11, The chemically crosslinked hydrogel network is a tissue engineering construct formed by covalently crosslinking macromonomers containing gelatin methacrylate (GelMA) and polyethylene glycol dimethacrylate (PEGDMA).

18. The structure according to claim 14, The salt ions are selected from Ca²⁺, Mg²⁺, Sr²⁺, Zn²⁺, Ti⁴⁺, or Al³⁺ in a tissue engineering construct.

19. A structure according to claim 14, The salt ions are tissue engineering constructs containing Ca²⁺.

20. The structure according to claim 14, The salt is a tissue engineering construct containing CaSO₄ or CaCl₂.

Images