Bioactive implants for the reconstruction of bone defects, bone deformities, and nonunions

Abstract

According to the present disclosure, a bioactive implant for orthopedic use is provided. The bioactive implant of the present disclosure includes a rod having a surface treated to modify the surface properties, a freeze-dried hydrophilic hydrogel network physically crosslinked to the surface of the rod via a charged polymer and salt ions, a biological material entrapped and carried by the hydrophilic hydrogel network, and a covalently bonded reactive macromonomer chemically crosslinked to the hydrophilic hydrogel network to enhance the physical crosslinking of the hydrophilic hydrogel network to the surface of the rod. The surface of the rod is coated with covalently bonded molecules, which are chemically crosslinked to the covalently bonded reactive macromonomer to enhance the adhesion of the hydrophilic hydrogel network chemically and physically crosslinked to the surface of the rod to the rod.

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 bioactive implants for the reconstruction of bone defects, bone deformities, and non-unions. [Background technology]

[0003] Segmental bone defects resulting from high-energy trauma, debridement procedures, or tumor resection remain a major challenge in the field of orthopedic surgery. Due to the combined functions of mechanical support and bone reconstruction, autogenous or allogeneic bone grafts have been regarded as the traditional methods for reconstructing segmental bone defects. However, these procedures usually reduce the patient's mobility, require multiple surgeries, and are only effective in reconstructing bone defects to a limited extent.

[0004] Modern limb salvage procedures for fracture fixation or reconstruction have been used to treat large bone defects for decades. In particular, the Ilizarov technique, also known as distraction osteogenesis (DO), is a relatively mature limb salvage procedure in the treatment of large bone defects. This procedure involves osteotomy followed by bone lengthening via an external or internal fixator. Depending on the type of treatment, the Ilizarov technique is divided into three categories: acute shortening and lengthening (monofocal approach), single-level bone transport (bifocal approach), and bilevel bone transport (trifocal approach). For the treatment of small bone defects (usually <2 cm), a monofocal approach is recommended to directly close the bone defect before bone lengthening. However, when the size of the bone defect is more than 3 cm, acute shortening may affect the soft tissues and damage the vasculature, resulting in limb ischemia. The bifocal approach uses an osteotomy at a site away from the defect site. The intersegmental segments are then displaced away from the cortical bone resection site and compressed at the defect site to maintain optimal bone length. In the trifocal approach, two distraction osteotomies are performed in addition to compression at the defect site. The bifocal or trifocal approach has shown some advantages in treating larger bone defects without limb discordance or soft tissue retraction. However, bone transport still has drawbacks, such as pin track infection, nonunion at the docking site, and inadequate bone consolidation. Among these drawbacks, nonunion at the docking site was observed in almost all patients who underwent bone transport. Inactive bone contact and soft tissue intrusion at the docking site tend to form a pseudoarthrosis. A secondary debridement procedure and bone grafting surgery are always essential to achieve final bone reconstruction at the docking site. In recent years, an increasing number of reports have addressed bone transport using metallic intramedullary nails (IM nails) as a method to shorten the external fixation period. However, there is no reliable evidence to suggest that metallic IM nails achieve early bone consolidation or reduce nonunion at the docking site. In addition, when metallic IM nails are used, additional surgery is required to remove the metallic IM nails.

[0005] Bone morphogenetic proteins (BMPs) are one of the most powerful bone inductive factors that play an important role in bone repair and reconstruction. In DO animal models, BMP-2, BMP-4, and BMP-7 have been shown to be highly expressed at bone reconstruction sites during the bone distraction phase and gradually decrease during the bone consolidation phase. Local injection of recombinant human (rh) BMP-2 or BMP-7 has been shown to promote bone formation in many DO models. The inventors hypothesized that by using a biodegradable intramedullary nail (IM nail), it would be possible to maintain the delivery of BMPs in a single surgery, promote bone healing, and reduce nonunion in bone transport without secondary surgery, grafting, or severe complications. The present invention solves at least some of the above problems. Summary of the Invention [Means for solving the problem]

[0006] In one embodiment, the present disclosure provides a method for reconstructing a bone defect, bone deformity, or bone nonunion. Also, the present disclosure provides a bioactive implant, the bioactive implant including a scaffold, a freeze-dried hydrogel network layer physically and chemically crosslinked to the scaffold and chemically bonded to the scaffold, and a biological material dispersed and supported within the freeze-dried hydrogel network layer. In one embodiment, the bioactive implant is coated.

[0007] The bioactive implant is sized to fit between two bone segments, into a bone tunnel, or at a fracture site, and is implanted between two bone segments, into a bone tunnel, or at a fracture site. When the bioactive implant is used as an intramedullary bioactive implant, the bioactive implant is sized to fit intramedullary implant between two bone segments. When the bioactive implant is used in a surgically created bone tunnel to treat foot, ankle, shoulder, hip, or joint injuries and disorders, the bioactive implant is sized to fit or slightly larger than a press fit into a surgically created bone tunnel at the anatomical location intended for treatment.

[0008] In another embodiment, the scaffold is an interconnected porous scaffold, the lyophilized hydrogel network layer is physically and chemically crosslinked and chemically bonded to the interconnected porous scaffold, and the biological material is dispersed and retained within the lyophilized hydrogel layer carried on the surface and within the pores of the interconnected porous scaffold.

[0009] In yet another embodiment, the present disclosure provides a bioactive implant for orthopedic surgery. The bioactive implant of the present disclosure includes a scaffold, a freeze-dried hydrogel network layer physically and chemically crosslinked to the scaffold and chemically bonded to the scaffold, and a biological material dispersed and supported within the freeze-dried hydrogel network layer. The bioactive implant has a size suitable for implantation between two bone segments or at a fracture site. When the bioactive implant is used as an intramedullary bioactive implant, the bioactive implant has a size suitable for intramedullary implantation between two bone segments. In one embodiment, the bioactive implant is coated.

[0010] In yet another embodiment, the scaffold is an interconnected porous scaffold, the lyophilized hydrogel network layer is physically and chemically bonded to the interconnected porous scaffold, and the biological material is distributed and supported on the surface and within the pores of the interconnected porous scaffold.

[0011] In yet another embodiment, the present disclosure provides a bioactive (intramedullary (IM)) implant device-based therapy for improved treatment of long bone defects, correction of bone deformity, or treatment of bone nonunion. The bioactive implant device includes a core scaffold and a bioactive hydrogel coating. The core scaffold is made of polymer, ceramic, metal, or composite. The core scaffold is porous or non-porous, and degradable or non-degradable. The hydrogel can be an interpenetrating network of physically crosslinked gels and covalently crosslinked gels. Examples of hydrogels include gelatin methacryloyl-alginate-based (GelMA-alginate) gels. Low molecular weight crosslinkers can be added to the hydrogel to increase crosslink density. Biological materials include growth factors and drugs. Growth factors include BMP-2, PDGF, IGF-1, FGF2. Growth factors are loaded into the hydrogel and exhibit a tunable or sustained release pattern. The bioactive implant device of the present disclosure can be used for distraction osteogenesis and as an adjunct therapy for other orthopedic injuries and disorders (FIGS. 1A-D). (1) Bone transport via IM implants; (2) Bone lengthening with IM implants (3) bone treatment with an IM implant or implant sized to fit into a bone tunnel for the treatment of a bone nonunion or bone defect; or (4) Treatment of bone defects with IM implants.

[0012] It should be noted that the bioactive implant device can be used in normal bone graft sites between bones and bone defects, not just in the intramedullary space, and is not limited to intramedullary applications. The scaffold (or IM nail) can be a solid rod or a porous scaffold, with metallic rods being load-bearing, among others.

[0013] The bioactive implant device of the present disclosure can effectively promote bone consolidation in bone distraction when patients undergo bone transport surgery, thereby preventing nonunion at the docking site. The bioactive implant device of the present disclosure can also effectively promote bone union in the treatment of nonunion. Figures 1A-D show the schematic design of the bioactive implant device of the present disclosure for bone healing under three different conditions.

[0014] In another embodiment, the present disclosure provides an implant device including a scaffold and a coating material. The core scaffold material is made of a polymer, ceramic, metal, or composite material, such as polycaprolactone-tricalcium phosphate (PCL-TCP) or polylactic-co-glycolic acid-TCP (PLGA / TCP). PCL-TCP filaments with a weight ratio of 80:20 (w / w) can be printed into a three-dimensional porous scaffold with pore sizes ranging from 100 μm to mm. The core scaffold material can be a metallic material such as titanium alloy, cobalt chromium alloy, magnesium alloy, zinc alloy, etc. The core scaffold can be porous or non-porous. The hydrogel coating layer can cover the whole or part of the core scaffold.

[0015] Regarding implantation, there are four different methods according to three applications (FIGS. 1A-D). (1) In the treatment of (large) bone defects, the implant device of the present disclosure is inserted between the distal and proximal bone segments through the gap of the bone defect during bone transport surgery or in a minimally invasive manner. The hydrogel coating layer should cover the length from the reconstruction site to the docking site. Both ends of the implant device are fixed by distal and proximal fixation pins. Of note, the pins for transporting the middle bone segment should not come into contact with the IM device. (2) In bone distraction, the implant device of the present disclosure is inserted into the long bone from anterior or posterior in a minimally invasive manner. The hydrogel coating layer should cover the length of the reconstruction site. The implant device is fixed by distal or proximal fixation pins. (3) In the treatment of nonunion, the implant device of the present disclosure is inserted into the long bone from anterior or posterior in a minimally invasive manner during a surgical debridement procedure. The hydrogel coating layer should cover the length of the nonunion gap. The implant device is fixed by distal or proximal fixation pins. (4) In the treatment of bone defects, the implant device of the present disclosure is inserted into the distal and proximal bone segments through the gap of the bone defect. The hydrogel coating layer should cover the length of the gap of the bone defect. The implant device is fixed by distal and proximal fixation pins, or internal or external fixation devices such as plates and IM nails.

[0016] An example of a biologically supported hydrogel coating is shown below. The hydrogel coating contains gelatin methacrylate (GelMA; 15%), alginate (1.25%), poly(ethylene glycol) dimethacrylate (PEGDMA; 2%), heparin methacrylate (HepMA; 1%), bone morphogenetic protein-2 (BMP-2; 200 μg / mL), and photoinitiator (0.3%) in deionized water. To synthesize GelMA macromonomer, gelatin was dissolved in deionized water (10% w / v) at 50 °C. To synthesize heparin methacrylate (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. The solution was reacted at room temperature for 1 h, and then 25 mg of APMA in 1 mL of MES was added and reacted at room temperature for 2 h. The HepMA solution was then 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). It was then lyophilized and stored at -80 °C.

[0017] An example of a bioactive implant device is shown below. For example, polycaprolactone-β-tricalcium phosphate (PCL-TCP) filaments carrying BMP-2-containing hydrogel were fabricated as shown in Figure 2A-P. PCL-TCP filaments with a weight ratio of 80:20 (w / w) and diameter of 0.9 mm were synthesized as described above, manually cut to produce 15 mm filaments, and immersed in 5N NaOH solution for 6 h. The filaments were subsequently washed three times with deionized water and incubated for 30 min at room temperature in MES buffer (100 mM) containing EDC (5 mg / mL) and NHS (5 mg / mL) to activate the surface carboxylic acid groups. The filaments 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 filaments were then washed three times with deionized water to remove unreacted GelMA and incubated in EDC / NHS (5 mg / mL) in MES buffer at room temperature for 15 min. The GelMA-coated filaments were then washed three times with deionized water and dried under vacuum. The GelMA-coated filaments were then incubated with CaSO in deionized water (100 mg / mL) at 60 °C. 4 The hydrogel-loaded filaments were immersed in the suspension and sonicated for 30 seconds. The filaments were then transferred to wells of a 24-well plate and dried under vacuum. The dried filaments were immersed in wells of a 96-well plate containing GelMA (15%), alginate (1.25%), PEGDMA (2%), HepMA (1%), protein (BMP-2, 200 μg / mL), and photoinitiator (0.3%) in deionized water for 2 minutes at 37°C. The hydrogel-loaded filaments were removed from the solution and placed in a dry well of another 96-well plate for 5 minutes. The hydrogel-loaded filaments were then irradiated with visible light for 15 minutes to covalently crosslink the GelMA, PEGDMA, and HepMA. The crosslinked hydrogel-loaded filaments were lyophilized and stored at -80°C.

[0018] An example of in vitro validation of a bioactive implant device to promote bone reconstruction using a rat bone transport model is shown in Figures 3A-H, 4A-C, 5A-C, 6A-F, and 7. Figures 3A-H show ex vivo and in vitro demonstration of bone transport with a customized metallic distraction osteogenesis fixator. Figures 4A-C show gross observations and union rates of femoral samples harvested on postoperative day 34 (POD34) or postoperative day 55 (POD55). Figures 5A-C show 3D reconstructed images and quantitative bone mass data of affected femurs on POD34 or POD55 as measured by micro-CT analysis. Figure 6 shows mechanical properties of affected femurs measured by three-point bending test on POD34 or POD55, normalized to their contralateral sections. Figure 7 shows representative histological results of affected femurs on POD34 or POD55.

[0019] The mechanical properties of the IMN+B6 group loaded with 6 μg BMP2 were 26.6% (P<0.001) in normalized maximum load, 11.6% (P<0.001) in normalized Young's modulus, and 31% (P<0.001) in normalized energy absorption when compared to the blank control (BLK) group, and 26.1% (P<0.001) in normalized maximum load, 11.3% (P<0.001) in normalized Young's modulus, and 30.5% (P<0.001) in normalized energy absorption when compared to the IMN group (Figure 6A-F). However, on POD34, no significant differences were found in the mechanical properties of the (IMN+B6) group when compared with the intramedullary nails loaded with 2 μg of BMP2 (IMN+B2) group (Figures 6A-F).

[0020] The histological data show longitudinal sections of the affected femur including the docking site and the reconstructed site (Figure 7). The results showed that in the BLK and IMN groups, soft tissue interposition was observed at the docking site on POD 34 and POD 55 (Figure 7). However, in the (IMN+B2) and (IMN+B6) groups, bone union was observed at the docking site on POD 34 and POD 55, without soft tissue interposition (Figure 7). In all groups, new bone was observed at the reconstructed site on POD 34 and POD 55 (Figure 7). Pin track infection was observed in some of the specimens in the BLK and IMN groups on POD 34 and POD 35 (Figure 7). In the (IMN+B2) group, cortical and trabecular bone were clearly observed not only at the docking site but also at the reconstructed site on POD 55, indicating enhanced bone remodeling (Figure 7). Furthermore, in the (IMN+B2) and (IMN+B6) groups, a very large amount of new bone was observed at the transplant site (FIG. 7).

[0021] The embodiments of the present disclosure can be applied as an intramedullary implant device for delivering growth factors for bone healing. An example is given below. Metal, polymer, or composite implants to deliver growth factors for the treatment of long bone defects as an adjunct to bone transport procedures; Metallic, polymeric or composite implants to deliver growth factors for the correction of bone deformities as an adjunct to distraction osteogenesis; Metal, polymer or composite implants for delivery of growth factors for the treatment or prevention of non-union; or Antibiotic-loaded metal, polymer, or composite implants for bone transport or bone distraction.

[0022] The embodiments of the present disclosure are beneficial in the field of segmental bone defects resulting from high-energy trauma, debridement procedures, or tumor resection, which remain a major challenge in the field of orthopedic surgery. Due to the combined functions of mechanical support and bone reconstruction, autologous or allogeneic bone grafts have been regarded as traditional methods for reconstructing segmental bone defects. However, these procedures usually reduce the patient's mobility, require multiple surgeries, and are only effective in reconstructing bone defects to a limited extent. Modern limb salvage surgery for fracture fixation or reconstruction has been used in the treatment of large bone defects for decades. In particular, the Ilizarov procedure, also known as distraction osteogenesis (DO), is a relatively mature limb salvage surgery in the treatment of large bone defects. This surgery involves osteotomy followed by bone lengthening via an external or internal fixator. Depending on the type of treatment, the Ilizarov procedure is divided into three categories: acute shortening and lengthening (monofocal approach), single-level bone transport (bifocal approach), and bilevel bone transport (trifocal approach). For the treatment of small bone defects (usually <2 cm), a monofocal approach with direct closure of the bone defect prior to distraction bone is recommended.

[0023] However, if the size of the bone defect is more than 3 cm, acute shortening may affect the soft tissue, damage the vasculature, and result in limb ischemia. The bifocal approach uses an osteotomy at a site away from the defect site. The middle interphalangeal segment is then displaced away from the cortical bone resection site and compressed at the defect site to maintain optimal bone length. The trifocal approach involves two distraction osteotomies in addition to compression of the defect. The bifocal or trifocal approach has shown some advantages in treating larger bone defects without limb discrepancy or soft tissue contraction. However, bone transport still has drawbacks, such as pin track infection, docking site nonunion, and inadequate bone consolidation. Among these drawbacks, docking site nonunion was observed in almost all patients who underwent bone transport. Inactive bone contact and soft tissue intrusion at the docking site tend to form a pseudoarthrosis. A secondary debridement procedure and bone grafting surgery are always essential to achieve final bone reconstruction at the docking site. In recent years, an increasing number of reports have addressed bone transport using metallic intramedullary nails (IM nails) as a method to shorten the duration of external fixation.

[0024] However, there is no reliable evidence to suggest that metallic IM nails achieve early bone consolidation or reduce nonunion at the docking site, and their use requires additional surgery to remove them.

[0025] The bioactive implant device according to the present disclosure was developed considering available regulation pathways, manufacturing process, distribution, and clinical applicability to address significant clinical challenges such as docking site nonunion and delayed bone consolidation. First, the bioactive implant device according to the present disclosure is based on FDA approved materials and growth factors and is applicable to rapid 510(K) regulation pathways. Second, the bioactive implant device according to the present disclosure is compatible with current surgical treatments and can be easily adapted by end-user orthopedic surgeons. Third, the bioactive implant device according to the present disclosure can be implanted in a single surgery required for osteotomy and fracture fixation for the treatment of long bone defects, bone deformities, or bone nonunions. The intramedullary device may be inserted at the proximal or distal end of the bone segment or may be inserted in a minimally invasive manner through a bone tunnel. The bioactive implant device according to the present disclosure can be fixed in situ by fixation pins. Fourth, no secondary surgery is required to remove tissue at the bone nonunion site and implant an allograft into the bone transport DO. During surgery for bone nonunion, no allograft is required. When made from biodegradable biomaterials, the bioactive implant device of the present disclosure is biodegradable. When made from metallic materials, the bioactive implant device of the present disclosure can provide mechanical support. Fifth, the storage and transportation conditions of the bioactive implant device of the present disclosure are good, which is convenient for manufacturers, distributors, and end-user surgeons.

[0026] The scaffold of the bioactive implant device of the present disclosure can be made of polymer, metal, or composite materials. The bioactive implant device of the present disclosure can be subjected to various surface treatments and then coated with hydrogels loaded with various doses of different growth factors depending on the application.

[0027] 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, which are bioactive implants. A schematic diagram of the method used to coat HyTEC constructs is shown in Figure 8. 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 surface of HyTEC, creating modified HyTEC (mHyTEC). The PCL-coated HyTEC can then be air-dried at 0-4°C. The physical properties of the mHyTEC constructs and the release rate of proteins can be tuned 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 9. After 14 days, 92% of the encapsulated BSA was released from uncoated BSA-loaded HyTECs, but when a protective PCL coating of one or three layers was formed using a PCL / acetone (10% wt / v) solution, the BSA release rate decreased to 92% after 70 days or 80% after 91 days (Figure 10). 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 coating were deposited (Figure 11).

[0028] In another characterization, the present invention can be described as follows: Distraction osteogenesis is one of the most successful surgical approaches for the treatment of large segmental bone defects. However, notable complications include prolonged bone sclerosis and nonunion at the docking site, especially when accompanied by bone transport. In the latter case, a secondary operation of bone grafting is always required to achieve bone bridging. Therefore, the inventors have fabricated an osteoinductive, biodegradable intramedullary (IM) nail by eluting bone morphogenetic protein-2 (BMP-2) from a biodegradable implant as an adjunct therapy to address the clinical challenges of bone transport described above. For proof of concept, the inventors fabricated an IM nail and tested it in a rat long bone transport model. First, polycaprolactone-tricalcium phosphate (PCL-TCP) filaments were coated with gelatin methacryloyl-alginate (GelMA-alginate) hydrogel loaded with BMP-2 (2 μg or 6 μg) to fabricate an IM nail. The release characteristics of BMP-2 in the IM nail were measured after freeze-drying and sterilization. The efficacy of the IM nail was then evaluated in a rat femoral bone transport model. The healing process was observed by X-ray every week. On the 34th or 55th postoperative day, samples were taken to macroscopically observe pin track infection and segmental bridging, and the mechanical properties, microstructure, and morphology of new bone formation at the docking site were evaluated. The results showed that BMP-2 was continuously released from the IM nail over a 21-day period. In the femoral bone transport model, pin track infection and insufficient bone consolidation were predominantly observed at both the docking site and the reconstruction site. The bone mass and three-point bending mechanical strength of the IM nail group containing BMP-2 were significantly greater than those of the surgical control group on POD34 and POD55. The IM nail containing BMP-2 also reduced pin track infection and promoted bone union at the docking site. Histological data also confirmed the superior effect of the IM nail containing BMP-2, with a higher mineral attachment rate in the BMP-2-containing IM nail group.No significant difference was observed in bone union effect between the two BMP-2 contents.Overall, the experiments demonstrated that the IM nail implant containing biodegradable BMP-2 significantly promoted bone consolidation, reduced pin infection, improved union rate at the docking site, enhanced load-bearing capacity, and led to early removal of the external fixator in distraction osteogenesis (DO) without the need for secondary surgery. This adjunctive therapy technology is expected to revitalize and revolutionize bone transport in the treatment of large bone defects and limb salvage surgery in the future.

[0029] According to the present disclosure, a method for reconstructing a bone defect, bone deformity, or bone nonunion is provided. The method of the present disclosure includes a step of preparing a bioactive orthopedic implant. The bioactive orthopedic implant includes: (i) a rod having a surface treated to modify the surface properties; (ii) a freeze-dried hydrophilic hydrogel network physically crosslinked to the surface of the rod via a charged polymer and salt ions; (iii) a biological material entrapped in the hydrophilic hydrogel network and thereby carried by the hydrogel network; and (iv) a covalently bonded reactive macromonomer chemically crosslinked to the hydrophilic hydrogel network to enhance the physical crosslinking of the hydrophilic hydrogel network to the surface of the rod. The method of the present disclosure also includes a step of implanting the bioactive orthopedic implant between two bone segments, in a bone tunnel, or at a fracture site. The bioactive orthopedic implant has a size suitable for implantation between two bone segments, in a bone tunnel, or at a fracture site. Alternatively, the bioactive orthopedic implant is sized to fit for intramedullary implantation between two bone segments, in a bone tunnel, or at a fracture site. The rod may be an interconnected porous rod, in which case the interconnected porous rod includes an embodiment of a central or internal channel in a porous scaffold / rod. The coating of the bioactive orthopedic implant may be applied to the entire or partial rod. Thus, different segments of the rod may also be coated differently. All of this depends on the type of reconstruction desired, as will be readily understood by those skilled in the art. In addition, the bioactive orthopedic implant may further include one or more (additional) coating layers. In a further embodiment, the rod is an interconnected porous rod, and the biological material is carried in the pores of the interconnected porous rod.

[0030] The present disclosure also provides a bioactive orthopedic implant. The bioactive orthopedic implant of the present disclosure includes (i) a rod having a surface treated to modify the surface properties, (ii) a freeze-dried hydrophilic hydrogel network physically crosslinked to the surface of the rod via a charged polymer and salt ions, (iii) a biological material entrapped in the hydrophilic hydrogel network and thereby carried by the hydrogel network, and (iv) a covalently reactive macromonomer that is chemically crosslinked to the hydrophilic hydrogel network to enhance the physical crosslinking of the hydrophilic hydrogel network to the surface of the rod. The surface of the bioactive orthopedic implant (rod) is coated with a covalent molecule, which is chemically crosslinked to the covalently reactive macromonomer to enhance the adhesion of the hydrophilic hydrogel network chemically and physically crosslinked to the surface of the rod to the rod. The rod is an interconnected porous rod, and the biological material is carried in the pores of the interconnected porous rod. The bioactive orthopedic implant is sized to be implanted between two bone segments, in a bone tunnel, or at a fracture site. Alternatively, the bioactive orthopedic implant is sized to be implanted intramedullary between two bone segments, in a bone tunnel, or at a fracture site. The coating of the bioactive orthopedic implant can be applied to the entire or partial rod. Thus, different segments of the rod can be coated with different coatings. All of this depends on the type of reconstruction desired, as can be easily understood by those skilled in the art. In addition, the bioactive orthopedic implant can further include one or more (additional) coating layers. [Brief description of the drawings]

[0031] If necessary, for further interpretation of the greyscales in the figures, the reader is referred to the priority documents for each figure.

[0032] [Figure 1A]1A-D show schematic diagrams of a bioactive implant device for bone healing under three conditions according to an exemplary embodiment of the present invention. "Blue" represents the hydrogel coating. "White dots" on the implant device represent the porous structure. FIG. 1A shows bone transport on the implant. Both ends of the implant are fixed with two fixation pins. [Figure 1B] FIG. 1B shows bone distraction via an implant. The implant can be inserted through a bone tunnel. One end of the implant is fixed with a fixation pin. [Figure 1C] Figure 1C shows a nonunion treated with an implant, both ends of which are fixed with two fixation pins. [Figure 1D] Figure ID shows a bone defect treated with an implant, both ends of which are fixed with fixation pins. [Figure 2A] 2A-P show the fabrication and characterization of an intramedullary (IM) nail implant according to an exemplary embodiment of the present invention. Figure 2A shows a schematic procedure for fabricating an osteoinductive, biodegradable IM nail implant. [Figure 2B] FIG. 2B shows an SEM image of the surface of the PCL-TCP filament after extrusion. [Figure 2C] FIG. 2C shows an SEM image of the surface of the PCL-TCP filament after extrusion. [Figure 2D] FIG. 2D shows an SEM image of the surface of PCL-TCP filaments after NaOH treatment and freeze / thaw. [Figure 2E] FIG. 2E shows an SEM image of the surface of PCL-TCP filaments after NaOH treatment and freeze / thaw. [Figure 2F] FIG. 2F shows an SEM image of the surface of a PCL-TCP filament after GelMA coating and ultrasonic treatment in a CaSO4 suspension. [Figure 2G] FIG. 2G shows an SEM image of the surface of a PCL-TCP filament after GelMA coating and ultrasonic treatment in a CaSO4 suspension. [Figure 2H]FIG. 2H shows an SEM image of the surface of the PCL-TCP filament after hydrogel coating. [Figure 2I] FIG. 2I shows an SEM image of the surface of the PCL-TCP filament after hydrogel coating. [Figure 2J] Figure 2J shows the effect of chemical treatment without freeze / thaw or sonication (chem; Bl), with freeze / thaw but without sonication (chem+FT; G), or with freeze / thaw and sonication (chem+FT+sonic; R) on the tensile modulus of PCL-TCP filaments. The control group was PCL-TCP filaments without any treatment (untreated, B). [Figure 2K] FIG. 2K shows the weights of PCL-TCP filaments without a hydrogel coating (rod; Br), filaments with a hydrogel coating but no BMP-2 (rod+gel; Bl), filaments with a hydrogel coating containing 2 μg BMP-2 (rod+gel / 2 μg BMP-2; G), and filaments with a hydrogel coating containing 6 μg BMP-2 (rod+gel / 6 μg BMP-2; R) (FIG. 2K). [Figure 2L] FIG. 2L shows the change in weight percentage of PCL-TCP filaments without hydrogel coating (rods; Br), filaments with hydrogel coating but no BMP-2 (rods+gel; Bl), filaments with hydrogel coating containing 2 μg BMP-2 (rods+gel / 2 μg BMP-2; G), and filaments with hydrogel coating containing 6 μg BMP-2 (rods+gel / 6 μg BMP-2; R). [Figure 2M] FIG. 2M shows a macroscopic view of an implant loaded with a wetted BMP-2-containing hydrogel. [Figure 2N] FIG. 2N shows a macroscopic view of an implant loaded with freeze-dried BMP-2-containing hydrogel. [Figure 2O]FIG. 2O shows the amount of BMP-2 released over a 28 day period from implants loaded with lyophilized hydrogels containing 2 μg (G) or 6 μg (R) of BMP-2. [Figure 2P] FIG. 2P shows the amount of BMP-2 released from implants loaded with preserved (Bl) or untreated (G) lyophilized hydrogels containing 2 μg of BMP-2 over 28 days. For the data in FIG. 2K and FIG. 2L, there were 8 samples per group. For the data in FIG. 2J, FIG. 2O and FIG. 2P, there were 3 samples per group. The data shown in FIG. 2J, FIG. 2O, FIG. 2P correspond to the mean ± SD. The data shown in FIG. 2K and FIG. 2L correspond to the mean and 95% confidence interval. [Figure 3A] 3A-H show ex vivo and in vivo demonstrations of bone transport with a customized metallic external fixator according to an exemplary embodiment of the present invention. Figure 3A shows the schematic design of a unilateral external fixator frame (longitudinal cross section or side view). [Figure 3B] FIG. 3B shows the final product of the external fixator frame and pins after fabrication. [Figure 3C] FIG. 3C shows a top view of the external fixator frame and pins secured to a rat femur after osteotomy, decorticectomy, and IM nail implantation. [Figure 3D] FIG. 3D shows the transfer of a bone slice through the IM nail from the proximal to the distal femur. [Figure 3E] FIG. 3E shows the transport of a bone slice through the IM nail from the proximal to the distal femur. [Figure 3F] FIG. 3F shows the transport of a bone slice through the IM nail from the proximal to the distal femur. [Figure 3G] FIG. 3G shows a close-up of the surgery after osteotomy and corticotomy. [Figure 3H] FIG. 3H shows a close-up view of the IM nail implant. [Figure 4A]4A-C show the macroscopic observations and union rates of femoral samples harvested on POD34 or POD55 according to an exemplary embodiment of the present invention. Animal groups included a blank control group (BLK), an IM nail alone group (IMN), an IM nail containing 2 μg of BMP-2 group (IMN+B2), and an IM nail containing 6 μg of BMP-2 group (IMN+B6). FIG. 4A shows the macroscopic view of the femoral sample. The white arrow points to the docking site. The two white dashed lines indicate the reconstruction site. [Figure 4B] FIG. 4B shows the bone union rate determined by macroscopic observation and micro-CT analysis. [Figure 4C] Figure 4C shows the rate of pin track infection assessed on POD 34 or POD 55 before sacrifice (n=8). Pin track infection was graded according to the Checketts-Otterburn classification. Only minor pin track infection was observed in some animals prior to sacrifice, including one or two loose pins in each animal. [Figure 5A] 5A-C show 3D reconstruction images and quantitative bone mass data of the affected femur on POD34 or POD55 measured by micro-CT analysis according to an exemplary embodiment of the present invention. Animal groups included a blank control group (BLK), an IM nail alone group (IMN), an IM nail containing 2 μg BMP-2 group (IMN+B2), and an IM nail containing 6 μg BMP-2 group (IMN+B6). FIG. 5A shows a representative 3D image (longitudinal half cut) of the affected femur. Two regions of interest (ROIs) of the docking site or reconstruction site were highlighted with a green dashed rectangle or a yellow dashed rectangle. [Figure 5B] FIG. 5B shows quantitative data of bone volume (bone volume / tissue volume, BV / TV) at the docking site or reconstruction site on POD34. [Figure 5C] Figure 5C shows quantitative data of bone volume (BV / TV) at the docking site or reconstruction site on POD55. Data are shown as normalized mean values ​​of BV / TV and SD (n=8 per time point). aP<0.05, aaP<0.01, aaaP<0.001, vs. BLK group; bP<0.05, bbP<0.01, bbbP<0.001, vs. IMN group. [Figure 6A] 6A-6F show mechanical properties measured by three-point bending test of affected femurs on POD34 or POD55, normalized with contralateral sections, according to an exemplary embodiment of the present invention. Animal groups include a blank control group (BLK), an IM nail alone group (IMN), an IM nail containing 2 μg BMP-2 group (IMN+B2), and an IM nail containing 6 μg BMP-2 group (IMN+B6). Mechanical parameters such as normalized maximum load rate, Young's modulus, and energy absorption rate are shown. FIG. 6A shows the mechanical property results on POD34. [Figure 6B] FIG. 6B shows the mechanical property results at POD34. [Figure 6C] FIG. 6C shows the mechanical property results at POD34. [Figure 6D] FIG. 6D shows the mechanical property results at POD55. [Figure 6E] FIG. 6E shows the mechanical property results at POD55. [Figure 6F] Figure 6F shows the results of mechanical properties on POD55. Data are shown as mean and standard deviation (n=8 per time point). aP<0.05, aaP<0.01, aaaP<0.001 vs. BLK group; bP<0.05, bbP<0.01, bbbP<0.001 vs. IMN group. [Figure 7] FIG. 7 shows representative histological results of affected femurs on POD34 or POD55 according to an exemplary embodiment of the present invention. Animal groups include a blank control group (BLK), an IM nail only group (IMN), an IM nail containing 2 μg BMP-2 group (IMN+B2), and an IM nail containing 6 μg BMP-2 group (IMN+B6). The red dashed rectangle indicates the docking site. The yellow dashed rectangle indicates the reconstruction site. The histology in this figure includes hematoxylin and eosin (H&E) and Masson's trichrome staining. [Figure 8] FIG. 8 shows a schematic diagram of the method used to coat a bioactive implant (a HyTEC construct), according to an exemplary embodiment of the present invention. [Figure 9]FIG. 9 shows images of mHyTEC with one layer of PCL coating (mHyTEC(1L)) or 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 10] FIG. 10 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 present invention. [Figure 11] FIG. 11 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 12] FIG. 12 shows in vivo dynamic X-ray images of affected femurs from POD5 to POD34 according to an exemplary embodiment of the present invention. Animal groups included a blank control group (BLK), an IM nail alone group (IMN), an IM nail containing 2 μg BMP-2 group (IMN+B2), and an IM nail containing 6 μg BMP-2 group (IMN+B6). White arrows indicate the bone defect site before bone transport on POD5. Red arrows indicate the bone defect site after bone transport on POD13-34 (after 3 weeks of consolidation). [Figure 13A]13A-13B show histomorphometry results of affected femurs on POD 34 according to an exemplary embodiment of the present invention. Animal groups included a blank control group (BLK), an IM nail alone group (IMIN), an IM nail containing 2 μg BMP-2 group (IMN+B2), and an IM nail containing 6 μg BMP-2 group (IMN+B6). The white arrows indicate the distance between the green and red fluorescent dyes, representing mineral deposition over the 10 days prior to sacrifice. [Figure 13B] Data are shown as the mean and standard deviation (n=3). ap<0.05, aap<0.01, aap<0.001, vs. BLK group; bp<0.05, bbp<0.01, bbp<0.001, vs. IMN group. [Figure 14] 14 shows the expression of osteocalcin (OCN), bone morphogenetic protein receptor II (BMPRII), or α-smooth muscle actin (α-SMA) in the periosteum of the docking site of the affected femur in POD according to an exemplary embodiment of the present invention. Animal groups included an IM nail only group (IN), an IM nail containing 2 μg BMP-2 group (IMN+B2), and an IM nail containing 6 μg BMP-2 group (IN+B6). White arrows indicate positive expression of osteogenic markers OCN, BMPRII, or periosteal stem cell marker α-SMA stained by immunofluorescence. [Figure 15A] Figures 15A-15F show coatings formed on the entire structure of a rod or implant or on selected sub-regions or structures, as desired, according to an exemplary embodiment of the present invention. Figures 15A-D are schematic diagrams of a rod or implant 1520 having a coating 1510. Figures 15E-F are schematic diagrams of a partially or fully coated rod-like implant or a 3D printed porous implant. Figure 15A shows the coating 1510 formed on one end of the implant 1520. [Figure 15B] FIG. 15B illustrates the case where coating 1510 is formed on a portion of implant 1520 other than the center. [Figure 15C] FIG. 15C illustrates the case where coating 1510 is formed only on the central portion of implant 1520. [Figure 15D] FIG. 15D illustrates the case where coating 1510 is formed over the entire surface of implant 1520. [Figure 15E] FIG. 15E shows a rod-shaped implant with a coating 1510 formed on one end. [Figure 15F] FIG. 15F illustrates a 3D printed implant 1530 with coating 1510 formed over the porous structure. [Figure 16] 16 shows time-lapse x-ray images of bone reconstruction in defects treated with Infuse bone grafts and HyTEC implants in a sheep metatarsal bone transport model, according to an exemplary embodiment of the present invention. POD indicates post-operative day. For example, POD37 indicates post-operative day 37. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0033] definition

[0034] 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.). Covalent 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 aminopropyl methacrylamide (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. An 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 in vitro. 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 reconstruction (regeneration) of hard or soft tissues such as bone, cartilage, tendons, ligaments, muscles, heart and heart valves.

[0035] The following illustrative description addresses the question of whether it is possible to promote bone healing and reduce nonunions in bone transport by introducing a biodegradable IM nail capable of sustained release of BMPs in a single operation without secondary surgery, implantation, or severe complications. The aim of the translational research and present invention is to solve the clinical challenges of prolonged or insufficient bone consolidation and high rates of nonunion in the treatment of large bone defects by a single treatment combining a clinically available surgical approach, distraction osteogenesis, with a novel bioengineered solution, an osteoinductive biodegradable IM implant device. To this end, the inventors first established a clinically relevant rat bone transport model by developing a customized external fixator. Second, the inventors developed a novel BMP-2-eluting biodegradable IM implant device made from US Food and Drug Administration (FDA) approved materials and growth factors, and investigated the sustained release efficiency ex vivo. Third, we investigated the effect of the novel BMP-2 eluting IM nail on bone consolidation and docking site union in an established rat bone transport model.

[0036] result

[0037] Synthesis and characterization of intramedullary nail implants

[0038] The fabrication procedure of osteoinductive biodegradable intramedullary nail implant is shown in Figure 2A. After extrusion of polycaprolactone-β-tricalcium phosphate (PCL-TCP, weight ratio 80:20) filaments (18.0 mm long, 1.1 mm diameter each), BMP-2-containing hydrogel was loaded onto the filament surface by three consecutive steps, as shown in Figure 2A. Scanning electron microscope (SEM) images of the surface of PCL-TCP filament corresponding to each step are shown in Figure 2B-I. As the results show, the filament surface after NaOH treatment and freezing / thawing (Figure 2D-E) had higher roughness and microporosity compared with the filament surface immediately after extrusion (Figure 2B-C). GelMA coating, and CaSO 4 After sonication in suspension, CaSO 4Microparticles were deposited on the surface of the filaments (Fig. 2F-G). A porous BMP-2-containing hydrogel layer with an average pore size of 8 μm was formed on the surface of the PCL-TCP filaments by surface-initiated physical crosslinking of alginate followed by covalent crosslinking of GelMA and PEGDMA (Fig. 2H-I). The effect of chemical treatment (chem) without freeze / thaw or sonication (NaOH treatment and GelMA coating), chemical treatment with freeze / thaw but without sonication (chem+FT), and chemical treatment with freeze / thaw and sonication (chem+FT+sonic) on the tensile modulus of PCL-TCP filaments is shown in Fig. 2J. No significant difference in tensile modulus was observed between any of the treated filament groups (chem, chem+FT, chem+FT+sonic) and the untreated filament group (untreated). The effect of hydrogel loading and BMP-2 loading on the implant weight is shown in Fig. 2K. The average weight of the filaments increased from 15.9 mg to 32.3 mg upon hydrogel loading. The average loading of hydrogel on the filaments was 112.5%, and no significant change was observed whether the hydrogel contained 2 or 6 μg of BMP-2 (Figure 2L). Representative images of implants loaded with wet or freeze-dried BMP-2-containing hydrogels are shown in Figures 2M-N, respectively. The hydrogel layer remained intact after freeze-drying (Figure 2N). The diameter of the filaments increased on average from 0.9 mm to 1.2 mm when coated with a freeze-dried hydrogel layer. The release kinetics of 2 μg or 6 μg of BMP-2 from freeze-dried hydrogel-loaded implants over a 28-day period is shown in Figure 2O. BMP-2 was released from the implants in a sustained manner over a 21-day period. The total amount of BMP-2 released over 21 days from implants loaded with 2 or 6 μg of BMP-2 was 1.6 μg (80%) or 4.3 μg (71.7%), with no significant change observed after day 21. At each time point, the amount of protein released from implants loaded with 6 μg of BMP-2 was significantly higher than that from implants loaded with 2 μg of BMP-2.To assess the shelf life of freeze-dried BMP-2-loaded implants, the release kinetics of BMP-2 from freeze-dried implants after 2 months of storage at 4°C was measured and compared with that from untreated freeze-dried implants (Figure 2P). No statistically significant difference was observed between the total amount of BMP-2 released from untreated and stored implants at any time point over the 28 days.

[0039] Ex vivo or in vitro implantation of intramedullary nails (IM nails) in bone transport

[0040] The performance of the external fixator was investigated ex vivo using femoral specimens from adult SD rats. First, based on the basic external fixator, a bone distraction frame was designed with two fixed ends and one movable part between them (Figure 3A). The metallic external fixator is composed of a unilateral frame and five fixation pins (Figure 3B). The fixator was then used to fix the proximal and distal ends of the femoral specimen, and bone slices were transported in either an antegrade or retrograde direction via the IM nail between the two ends (Figure 3C-F). Five pins were inserted into the lateral site of the femur by a customized drill guide. An 8 mm osteotomy was performed in the femoral shaft. A 4 mm cortical incision was also made for bone transport. The 4 mm bone slice was then fixed with fixation pins inserted into the first cortical layer. The IM nail was inserted into the medullary canal, and both ends of the IM nail were fixed with two adjacent fixation pins, penetrating the medullary canal of a 4 mm bone slice. Note that the middle pin was inserted without contacting the IM nail in the medullary canal. During surgery, IM nails with or without BMP-2 coating were implanted into the animals (Figure 3G-H). After surgery, the animals were subjected to DO treatment. The treatment period of DO treatment consisted of three stages: a waiting period of 5 days, an extension period of 8 days, and a consolidation period of 21 days (3 weeks) or 42 days (6 weeks).

[0041] From the walking ability recorded on POD55 (after a 6-week hardening period), we found that animals implanted with IM nails that did not contain BMP-2 had poor walking ability and even showed paralytic gait after removal of the fixator one week earlier (POD48). Surprisingly, when IM nails coated with BMP-2 (2 μg) were implanted and the fixator was removed on POD48, the animals moved quickly and smoothly. On POD34 or POD55, femoral samples connected to the fixator were carefully collected and stored for further evaluation (Figure 4A).

[0042] The samples were macroscopically observed after harvesting. As shown in Figure 4A-B, the majority of femoral samples showed nonunion at the docking site in the blank control (BLK) and IM nail only (IMN) groups, while the union rate increased from 12.5% ​​on POD 34 to 25% on POD 5 in the BLK group, and from 0% on POD 34 to 37.5% on POD 55 in the IMN group. In addition, nonunion was observed at the docking site in some samples in the BLK and IMN groups (Figure 4A). Surprisingly, in all femoral samples, the IM nail group containing 2 μg BMP-2 (IMIN+B2, POD 34 or POD 55) and the IM nail group containing 6 μg BMP-2 (IMN+B6, POD 34) showed 100% union at the docking site (Figure 4A-B). It was also found that the reconstruction site after bone transport in all samples achieved 100% union. The union rate of the specimens was confirmed in duplicate by micro-computed tomography (CT) analysis.

[0043] After a 3- or 6-week curing period, animals showed minor infections at the pin sites, but no major infections were observed (Figure 4C). However, we found that when IM nails without BMP-2 coating (containing no BMP-2) were implanted, many animals had minor pin track infections. Pin track infections were measured during the fixator dismantling process before sacrifice. Although some animals showed pin loosening, external fixation could be continued for the duration of the study. As a result, pin track infection rates of 25% on POD34 or 50% on POD55 were observed in the BLK group, and 37.5% on POD34 or 50% on POD55 in the IMN group (Figure 4C). Interestingly, when IM nails containing BMP-2 were implanted, no infections occurred, regardless of the BMP-2 content or sacrifice time point (Figure 4C).

[0044] 2D and 3D radiological assessment of bone healing

[0045] Dynamic X-ray imaging was performed to observe the bone healing process. As shown in Figure 12, after a 5-day waiting period (POD5), an 8 mm bone defect was observed in each animal. After an 8-day extension period (POD13), an 8 mm bone defect was observed in the other sites of each animal. After a one-week consolidation period (POD20), callus formation was observed at the defect sites in the (IMIN+B2) and (IMN+B6) groups (Figure 12). The mineral density of the callus in the two groups containing BMP-2 gradually increased throughout the consolidation period (POD20-POD12) (Figure 12). Meanwhile, only limited callus formation was observed in the blank control group or the group with only the IM nail until POD34 (Figure 12).

[0046] Micro-CT analysis was performed to reconstruct 3D images of the affected femur and to measure the quantitative bone volume at two regions of interest (ROIs), i.e., the bone defect site and the docking site, on POD34 or POD55. As the results showed, nonunion was observed at the docking site on POD34 and POD55 in the BLK and IMN groups (Figure 5A-C). In contrast, obvious bone fusion and a significant increase in bone volume were observed at the docking site in the (IMN+B2) group (normalized BV / TV +30.1%, P<0.001 on POD34; normalized BV / TV +53.4%, P<0.05 on POD55) and (IMIN+B6) groups (normalized BV / TV +35.2%, P<0.001 on POD34) compared with the IMN group (Figure 5B-C). Compared with the BLK group, the (IMN+B2) group showed a significant increase in normalized BV / TV of 36.8% (P<0.001) or 30.5% (P<0.001) on POD34 or POD55, and the (IMN+B6) group showed a significant increase in normalized BV / TV of 41.9% (P<0.001) or 57.0% (P<0.001) on POD34 or POD55 (Figure 5B-C). In addition, new bone was observed at the reconstruction site in all groups (Figure 5A). A significant increase in bone volume was observed at the reconstruction site in the (IMN+B2) group (+16.1% in normalized BV / TV, P<0.05 at POD34; +26.1% in normalized BV / TV, P<0.05 at POD55) and (IMN+B6) group (+13.2% in normalized BV / TV, P<0.05 at POD34) compared with the IMN group (Figure 5B-C). At the reconstruction site, no significant difference was observed between the BLK group and the BMP-2-containing group on POD34 (Figure 5B). Also, no significant difference was observed in bone volume in the two ROIs between the two BMP-2-containing IM nail groups (Figure 5B).

[0047] Mechanical properties of affected femurs

[0048] Three-point bending mechanical testing was performed to measure the mechanical properties of the femoral specimens (Figure 6A). Parameters such as maximum load rate, Young's modulus, and energy absorption rate of the affected femur were normalized by those of the contralateral intact control bone. Due to the high rate of nonunion in the BLK and IMN groups, the number of femoral specimens available to complete all mechanical tests was very limited. For nonunion specimens, mechanical parameters were set to 0 as test failure. As the results show, except for the increased energy absorption rate in the IMN group at POD55, no significant differences were observed in other mechanical parameters between the BLK and IM nail groups (Figure 6A-F). In addition, the mechanical properties of the (IMN+B2) group were significantly improved compared to the BLK group, with the normalized maximum load rate, Young's modulus, and energy absorption rate increasing by 22.7% (P<0.01), 7.6% (P<0.05), and 18.6% (P<0.05), respectively, on POD34 (Figures 6A-C), and by 34.1% (P<0.001), 30.4% (P<0.01), and 18.9% (P<0.001), respectively, on POD55 (Figures 6D-F). In addition, the mechanical properties of the (IMIN+B2) group were also significantly improved compared to the IM group, with the normalized maximum load rate, Young's modulus, and energy absorption rate increasing by 22.2% (P<0.01), 7.3% (P<0.05), and 20.1% (P<0.05), respectively, on POD34 (Figures 6A-C) and by 29.4% (P<0.001), 31.0% (P<0.01), and 7.5% (P>0.05), respectively, on POD55 (Figures 6D-F).

[0049] The mechanical properties were also significantly improved in the (IMN+B6) group, with the normalized maximum load rate, Young's modulus, and energy absorption rate increasing by 26.6% (P<0.001), 11.6% (P<0.001), and 31% (P<0.001) compared to the BLK group on POD34, and by 26.1% (P<0.001), 11.3% (P<0.001), and 30.5% (P<0.001) compared to the IMN group (Figures 6A-C). However, no significant differences in mechanical properties were observed when comparing the (IMN+B6) and (IMN+B2) groups on POD43 (Figures 6A-C).

[0050] Histological analysis

[0051] The histological data shows longitudinal sections of the affected femur including the docking site and the reconstructed site (Figure 7). The results showed evidence of soft tissue interposition at the docking site in the BLK and IMN groups on POD34 and POD55 (Figure 7). However, in the (IMN+B2) and (IMN+B6) groups, bone union was achieved at the docking site on POD34 and POD55, and no soft tissue interposition was observed (Figure 7). New bone was observed at the reconstructed site in all groups on POD34 and POD55 (Figure 7). Pin track infection was observed in some specimens in the BLK and IMN groups on POD34 and POD55 (Figure 7). In the (IMIN+B2) group, cortical and trabecular bone were clearly identified at the docking site as well as the reconstructed site on POD55, indicating that bone remodeling was promoted in the animals (Figure 7). Furthermore, in the (IMN+B2) and (IMN+B6) groups, a very large amount of new bone was observed at the implantation site (FIG. 7).

[0052] Histological morphometry was performed to measure the dynamic bone mineralization rate in the docking and reconstruction sites of the affected femur on POD 34. Images and quantitative data from in vitro fluorescent labeling showed that the mineralization rate (MAR) was significantly increased by 128.1% (P<0.001) and 126.0% (P<0.001) in the reconstruction sites and 21.0% (P>0.05) and 27.0% (P<0.05) in the docking sites in (IMIN+B2) and (IMN+B6) compared with the BLK group (Figures 13A-B). Similarly, the mineral apposition rate (MAR) in (IMN+B2) and (MN+B6) was significantly increased by 142.6% (P<0.003) and 140.4% (P<0.001) at the docking site and by 33.0% (P<0.05) and 39.6% (P<0.01) at the reconstruction site compared to the IMN group (Figures 13A-B).

[0053] To understand the underlying mechanism, immunofluorescence staining and immunohistochemistry staining were performed. Figure 14 shows that high expression levels of bone formation marker (osteocalcin; OCN), BMP-2 receptor II (BMPRI; BI), and periosteal stem cell marker (α-SMA) were observed in the (IMN+B2) and (IMN+B) groups. These data indicate that bone formation is active at the docking site and that periosteal stem cells are involved in the healing process of the docking site. The high expression level of BMPRII promotes the binding of BMP-2 to osteogenic progenitor cells.

[0054] Consideration

[0055] Long bone transport DO model

[0056] We have successfully established a novel long bone transport DO model in rats, which mimics the clinical environment and outcomes of human patients undergoing bone transport surgery, who generally show normal bone reconstruction at the reconstruction site but have a high prevalence of nonunion and pin track infection at the docking site. Previous studies have reported long bone transport models using relatively large animals such as sheep and rabbits. Bone transport models in dogs and rabbits are also found in the field of mandibular tissue reconstruction. To the best of our knowledge, this is the first report of a rat bone transport model. Rodent models are traditionally used as the first choice for in vitro testing in orthopedic research due to advantages such as ease of availability and cost-effectiveness, before subsequent experiments and clinical studies in larger animal models. However, due to their small body size, rodents have rarely been used in traditional bone transport experiments, which require specialized hardware and specialized surgical expertise that allow for surgery in limited spaces and long-term extension. The present inventors have custom designed and manufactured a unique novel external fixator and drill guide for rat long bone surgery to facilitate the experimentation of the present study.

[0057] IM nail as a carrier for BMP-2 delivery

[0058] Wasserstein I is considered one of the pioneers of IM implants and their combination with an external fixator. He has been using this technique since 1963. However, secondary long bone allografts were required immediately after distraction osteogenesis. Paley first introduced the concept of combining an external fixator with an IM nail in situ in 1997 to shorten the duration of the external fixator. IM nails are usually made of stainless steel or titanium alloys. The advantage of such metallic IM nails is the high patient satisfaction rate, which is in part due to fewer psychosocial concerns. The disadvantages are the additional cost of the IM nail and the need for removal of the IM nail after osteoconsolidation.

[0059] As early as 1992, biodegradable IM nails made from copolymers of polyglycolic acid and polylactic acid were developed for intramedullary fixation. No significant difference in bone healing was observed between the control group, which used Kirschner wires for fixation of extra-articular fractures, and the biodegradable IM nail group. Since open wounds are exposed to bacteria and other infectious microorganisms, various types of antibiotic-eluting IM nails have been developed for fracture fixation and other bone reconstruction procedures. Recent clinical studies have shown that metallic IM nails with antibiotic-containing polymethylmethacrylate cement cores have less infection, faster bone consolidation, and fewer complications compared with standard nails in treating open tibial fractures. However, the inventors have not found such reports in either patient or animal models using biodegradable IM nails for bone transport. In the present invention, a novel technique for loading a BMP-2-containing hydrogel layer on PCL-TCP filaments was used to fabricate osteoinductive biodegradable IM nail implants. PCL-TCP composites have biocompatibility, mechanical stability, bioresorbability, and osteoconductivity. However, PCL-TCP constructs lack osteoinductive factors that induce bone formation and promote bone healing.

[0060] Surface coatings have been used to immobilize proteins on the surface of scaffolds for tissue engineering applications. However, the amount of protein that can be loaded on the surface of scaffolds by surface coating is generally low, and the release rate of the protein is also fast. For example, the loading amount of bovine serum albumin on the surface of hydroxyapatite-based scaffolds coated with chitosan and sodium hyaluronate by layer-by-layer (LBL) technique was lower and the release rate was faster than that of uncoated scaffolds. Hydrogels with high water content and porous microstructure provide a platform for sufficient loading and sustained release of proteins. However, loading soft hydrogels on the surface of rigid constructs is difficult due to the mismatch of mechanical properties at the interface. In addition, scaffolds loaded with protein-containing hydrogels using conventional techniques had to be used immediately after loading the hydrogel to avoid water evaporation from the hydrogel and protein denaturation. In the disclosed IM nail implant, the BMP-2-containing hydrogel layer is heavier than the PCL-ICP filament itself and is completely integrated with the rigid filament surface. The PCL-TCP filament with a weight ratio of 80:20 maintained a rigid filament shape as the core of the IM nail and remained stable when coupled with a fixation pin in both ex vivo and in vitro tests. After fabricating the PCL-TCP filament by extrusion, the filament was successively treated to increase hydrophilicity, improve hydrogel adhesion, and promote surface-initiated crosslinking. NaOH treatment and freeze / thawing impart hydrophilicity to the polyester surface by cleavage of ester bonds to carboxyl and hydroxyl groups and formation of micropores on the surface, respectively. Coating the surface with GelMA formed double bonds on the surface for covalent bonding with hydrogel. At a temperature (60°C) close to the melting point (S) of PCL, CaSO 4 By sonication in suspension, CaSO microparticles were deposited and supported on the soft surface of the GelMA-modified filaments, and none of the treatment steps had a detrimental effect on the tensile stiffness of the filaments. 4When the scaffold treated with PEG 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 GelMA and PEGDMA macromonomers in the physically crosslinked hydrogel were covalently crosslinked in the next step to form a robust interpenetrating network. In addition, because heparin has a high affinity for BMP-2, HeMA was incorporated into the hydrogel to extend the release of BMP-2. The addition of HeMA to an alginate-based hydrogel was shown to extend the release period of BMP-2 and improve subcutaneous bone formation in mice. The hydrogel loaded on the PCL-TCP filaments was heavier than the filaments themselves. Thus, the disclosed method can be used to load a sufficient amount of BMP-2 to promote bone remodeling in vivo. Furthermore, the hydrogel network remained integrated with the PCL-TCP filaments even after freeze-drying. Release data from freeze-dried BMP-2-loaded implants loaded with 2 μg or 6 μg of BMP-2 showed that 80.0% or 71.7% of the initially loaded BMP-2 was released after 21 days. Storage of the BMP-2-loaded implants for 2 months did not affect the activity and release kinetics of BMP-2. Thus, the disclosed method can be used to fabricate storable osteoinductive biodegradable intramedullary nail implants.

[0061] From a clinical perspective, we have successfully fabricated biodegradable IM nails using materials approved for clinical applications by the US Food and Drug Administration (FDA), such as PCL, TCP, alginate, bioactive factor th-BMP-2, and gelatin for food processing, and these combinations may form a promising medical device without many obstacles in future clinical applications. Among these materials, PCL-TCP composites have attracted widespread attention in bone tissue engineering due to their high biocompatibility, long-term degradability, suitable mechanical properties, and osteoconductivity. In the present invention, the PCL-TCP composite with a weight ratio of 80:20 maintained a rigid filament shape as the core of the IM nail and maintained stability when coupled with a fixation pin in both ex vivo and in vitro tests. GelMA / alginate composite hydrogels are also frequently applied in drug and cell delivery due to their adjustable mechanical properties and excellent biocompatibility. As the results show, due to the excellent integration with the PCL-TCP core, the GelMA / alginate shell formed after freeze-drying has the ability to sustain the sustained release of rhBMP-2 for 28 days with the assistance of heparin, which helps to protect BMP-2 from degradation and maintain the sustained release properties. In addition, the biodegradable IM nail showed excellent biocompatibility in animals, and showed almost the same rate of pin track infection and nonunion at the docking site as the blank control group, indicating a safe approach of this IM nail implantation. More and more studies have revealed that periosteal stem cells play a major role in DO-assisted bone defect healing. Although IM nails may endanger the bone marrow by occupying the medullary canal after implantation, in this study, bone healing is maintained when the periosteum remains intact on the bone segment. Evidence from clinical studies has already shown that metallic IM nails exert mechanical support during bone transport procedures without delaying bone healing.

[0062] BMP-2 and BMP-7 have been approved for clinical use in open fractures of long bones, nonunions, and spinal fusion. However, BMPs need to be administered appropriately to achieve satisfactory clinical outcomes and avoid potential side effects. The main role of the BMP delivery system is to retain growth factors at the bone injury site for a long time, as well as provide optimal biodegradability and mechanical support for tissue ingrowth. A prospective, randomized, controlled, single-blind clinical trial in 450 patients revealed that BMP-2 implants (BMP-2 supported on an absorbable collagen carrier) were superior to standard treatments (intramedullary nailing and routine soft tissue management) in reducing secondary interventions, promoting fracture and wound healing, and reducing infection rates in patients with open tibial fractures. BMP-2 (25–750 μg per rabbit) or BMP-7 (20 μg per rat) were applied locally by injection or sponge carrier to promote bone consolidation in rabbit or rat distraction osteogenesis models. Densitometric results showed a dose-dependent effect.

[0063] The present invention further demonstrated that even with the incorporation of a very low dose (2 μg per rat) of BMP-2, the IM nail was effective in promoting bone healing and allowing early removal of the fixator as early as POD 48 (after a 5-week bone consolidation period). This is the first report of specific delivery of BMP-2 by intramedullary delivery via a biodegradable IM nail, which can be adapted to current clinical practice, minimizing interruptions to DO procedures while immobilizing the effect of promoting localized bone healing.

[0064] Bone sclerosis at the docking site and reconstruction site

[0065] Nonunion at the docking site has been recognized as a frequent problem in bone transport, resulting in substantial prolongation of the healing process, secondary surgery for bone grafting, and delayed fixator removal. Nonunion at the docking site is believed to be due to inactive bone contact and significant soft tissue interposition. Once bone transport is complete, the hematoma is gradually replaced by fibrocartilage tissue, as a cavity exists before the transported bone segment contacts the target segment. The distal end of the transported segment and the docking end of the target segment are also sealed by fibrocartilage tissue. Macroscopic observation and imaging results showed that the bone transport animal model, with or without the use of IM nail implants, had a very low union rate at the docking site (0-37.5%) and even pseudoarthrosis, which was consistent with conventional clinical outcomes. These results also indicate that the IM nail itself does not affect bone consolidation at the docking site. In the nonunion specimens, soft tissue interposition or even pseudoarthrosis at the docking site was observed.

[0066] As mentioned above, bone consolidation at the docking site is the rate-limiting step in the entire healing process. Strategies to improve bone consolidation at the docking site have focused on surgical procedures such as acute shortening, bone grafting, compression, alternating compression-distraction, or bone marrow transplantation combined with demineralized bone matrix. The exciting results obtained in this disclosure show that all animals treated with IM nails incorporating 2 μg or 6 μg of BMP-2 achieved union at the docking site as early as POD 34, and mechanical properties improved at both time points without secondary surgery or grafting. Immunofluorescence and immunohistochemistry results revealed that osteogenic marker (OCN), periosteal stem cell marker (α-SMA), and BMPRII were highly expressed in the periosteum at the docking site. We hypothesized that BMP-2 released from the IM nail would have a positive effect on the docking site by maintaining active bone formation at the distal and docking ends of the segments and promoting bone fusion where the two ends meet. Such sustained BMP-2-eluting activated bone formation may also be useful in preventing soft tissue invasion into the docking site and converting migrating fibroblasts into bone tissue. The inventors hypothesized that BMP-2 released from the IM nail implant binds to its receptor, BMPRIl, and significantly promotes the migration and bone formation of periosteal stem cells present at both ends of the docking site. However, further research is needed to determine the mechanism by which sustained BMP-2 release at the docking site works.

[0067] In addition to sufficient bone consolidation at the docking site, the BMP-2-containing group also achieved early bone consolidation at the reconstruction site. This was indicated by a higher BV / TV ratio analyzed by micro-CT and mechanical testing at POD55. Although mechanical testing data showed that mechanical properties were not fully restored, the walking ability of the animals in the BMP-2-containing group returned to normal after removal of the fixator 1 week earlier (POD48), indicating that the reconstruction site was sufficiently mature (consolidated) to withstand the weight of daily activities. Accelerated bone consolidation was attributed to the sustained release of BMP-2 from the IM nail into the expanded, vascularized soft reconstruction, with a constant release kinetic controlled by the ratio of the amount of extension to the increasing amount of BMP-2 exposure along the IM nail. The bone consolidation period varies greatly depending on local and systemic health. However, in clinical practice, conventional rates were reported after a 1-week bone consolidation period for 1 mm of distraction osteogenesis. The results of the rodent model showed that the IM nail containing BMP-2 can achieve a high bone consolidation rate of 1.6 mm per week, even when considering the species difference. It is also very encouraging that the changes in BV / TV measured by micro-CT and MAR measured by histomorphometry at the reconstruction site in the BMP-2-containing group were not significant at the early time point (POD34) but were significant after the bone consolidation period (POD55) compared to the BLK or IMN groups. These results are important because they suggest that the BMP-2-containing IM nail implant of the present disclosure did not produce early bone consolidation that would prevent further lengthening in the treatment of clinically important large human bone defects.

[0068] Clinical indications and future prospects

[0069] The presented results and embodiments demonstrate that the novel biodegradable BMP-2-containing IM nail implant of the present disclosure safely and efficiently promotes bone consolidation at both the reconstruction and docking sites in DO procedures, and facilitates early removal of the external fixation device without secondary surgery. With such promising results and most of the materials already approved for medical use by the FDA, the inventors believe that the novel biodegradable BMP-2-containing IM nail of the present disclosure has great potential for clinical application in combination with bone transport technology.

[0070] IM nail implants can also be designed for in vivo studies in large animals or clinical trials, where the combination of a PCL-TCP porous scaffold and a GelMA-alginate-BMP-2-containing hydrogel coating with hybrid 3D printing technology provides sufficient mechanical support, increases BMP-2 loading, and facilitates vascularization after implantation. We expect that patients undergoing bone transport and limb lengthening will greatly benefit from this single-surgery approach, where external fixators can be removed earlier, secondary surgeries avoided, and pin track infections significantly reduced.

[0071] Materials and Methods

[0072] Study design

[0073] Biodegradable implants were fabricated as osteoinductive IM nail (IMN) grafts to guide the reconstruction of 8 mm femoral segment defects in bone transport DO. The disclosed biodegradable implants consist of a composite of PCL / TCP (4 / 1) filaments coated with a freeze-dried hydrogel containing GelMA (15%), alginate (1.25%), PEGDMA (2%), HepMA (1%), and photoinitiator (0.3%) in deionized water supplemented with BMP-2 (2 μg or 6 μg). The surface morphology, tensile strength, and in vitro protein release kinetics of the disclosed biodegradable implants were characterized. The disclosed biodegradable implants were implanted into the medullary canal of a rat transport DO model. The DO protocol in this study consisted of a 5-day waiting period, an 8-day extension period, and a 21-day (PO34D) or 42-day (POD55) hardening period. X-ray imaging, micro-CT analysis, mechanical testing, histological testing, and histomorphometry were performed to evaluate the effect of the IM nail device on bone defect healing following distraction osteogenesis (DO).

[0074] chemicals

[0075] 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. 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), type A gelatin, heparin, and calcium sulfate dihydrate (CaSO 4) was purchased from Sigma-Aldrich. 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 BMP-2 protein was provided by Medtronic. Human BMP-2 ELISA kit was purchased from Sigma-Aldrich.

[0076] Synthesis of PCL-TCP filaments

[0077] PCL-TCP filaments with a diameter of 0.9 mm and a weight ratio of PCL to TCP of 80:20 were synthesized. Briefly, 80 g of PCL and 20 g of TCP were dissolved separately in 800 mL and 400 mL of DMF, respectively, and stirred at 80 °C for 3 h. Subsequently, the PCL and TCP solutions were mixed and stirred for 1 h. The mixture was then precipitated in 4 L of water to fabricate PCL-TCP composite sheets. The PCL-TCP composite sheets were washed with water, and the residual solvent was evaporated in a fume hood at ambient temperature for 24 h. The dried PCL-TCP composites were cut into pellets and extruded using a home-made screw extruder.

[0078] Synthesis of gelatin methacrylate and heparin methacrylate.

[0079] To synthesize the GelMA macromonomer, gelatin was dissolved in deionized water (10% w / v) at 50 °C. To the gelatin solution, methacrylic anhydride was added 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.

[0080] To synthesize heparin methacrylate (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. After reacting for 1 h at room temperature, 25 mg of APMA in 1 mL of MES was added to the solution and reacted for 2 h at room temperature. The HepMA solution was then dialyzed against deionized water at ambient temperature for 3 days using dialysis tubing with a molecular weight cut-off of 6–8 kDa (Spectrum Laboratories, Rancho Dominquez, CA, USA), and then lyophilized and stored at -80 °C.

[0081] Coating of PCL-TCP filaments with BMP-2-containing hydrogel

[0082] The procedure for coating PCL-TCP filaments with BMP-2-containing hydrogel is shown diagrammatically in Figure 2A. PCL-TCP filaments with a diameter of 0.9 mm were synthesized, manually cut into 18 mm filaments, and immersed in 5N NaOH solution for 6 h. The filaments were then washed three times with deionized water and then incubated at room temperature for 30 min in MES buffer (100 mM) containing EDC (5 mg / mL) and NHS (5 mg / mL) to activate the carboxylic acid groups on the filament surface. Next, the filaments were washed three times with deionized water and then incubated at 37 °C for 1 h in a 2% solution of GelMA in MES buffer. The filaments were then washed three times with deionized water to remove unreacted GelMA, and then incubated at room temperature for 15 min in EDC / NHS (5 mg / mL) in MES buffer. The GelMA-coated filaments were then washed three times with deionized water and then dried under vacuum. The GelMA-coated filament was then dissolved in CaSO in deionized water (100 mg / mL) at 60 °C. 4The hydrogel-loaded filaments were immersed in the suspension and sonicated for 30 seconds. The filaments were then transferred to wells of a 24-well plate and dried under vacuum. The dried filaments were immersed for 2 minutes in wells of a 96-well plate containing GelMA (15%), alginate (1.25%), PEGDMA (2%), HepMA (1%), protein (BMP-2: 200 μg / mL), and photoinitiator (0.3%) in deionized water at 37 °C. The hydrogel-loaded filaments were removed from the solution and left in the dried wells of another 96-well plate for 5 minutes. The hydrogel-loaded filaments were then irradiated with visible light for 15 minutes to covalently crosslink GelMA, PEGDMA, and HepMA. The hydrogel-loaded filaments were then stored at -80 °C and lyophilized.

[0083] Characterization of tensile modulus, hydrogel loading, and surface morphology

[0084] The tensile modulus (tensile modulus) of the PCL-TCP filaments was tested using an Instron 5944 uniaxial testing system (Instron Corporation, Norwood, Massachusetts, USA) with a 2 kN load cell and a 1 N preload. Tensile modulus measurements were performed at a displacement rate of 1% strain / sec up to 25% strain. The slope of the linear region of the stress vs. strain curve was taken as the tensile modulus. Five samples per group were used for tensile modulus measurements.

[0085] The hydrogel loading rate (%) was calculated from the scaffold weight before hydrogel loading (Wb) and the scaffold weight after hydrogel loading (Wa) using the following formula: Eight samples per group were used for measuring the hydrogel loading rate. Hydrogel loading rate (%) = 100 × (Wa-Wb) / Wb

[0086] To visualize the surface morphology, the filaments were immersed in liquid nitrogen and freeze-dried. 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.

[0087] Protein release

[0088] To measure the release kinetics, filaments coated with hydrogels containing 2 μg or 6 μg BMP-2 were freeze-dried and incubated in 1 mL of PBS at 37 °C for 28 days. At each time point, the amount of BMP-2 in the release medium was measured using ELISA, and the release medium was replaced with fresh PBS. To investigate the effect of storage of the implants on the activity and release kinetics of the protein, filaments coated with hydrogels containing BMP-2 were stored at 4 °C for 2 months and the release kinetics of BMP-2 from the stored implants was measured and compared to untreated implants.

[0089] Design of an external fixator for bone transport

[0090] A unilateral external fixator was specially designed and customized for bone transport in the Sprague-Dawley rat model. The unilateral external fixator is composed of two parts: one frame (32 mm long) and five fixation pins (1.2 mm diameter, 22 mm long). The frame has two fixed ends, each fixed by two fixation pins, and is used to fix the proximal and distal bone segments after osteotomy. The frame also has one movable part between the two fixed ends, which is used to lock the last pin to fix the movable bone slice after cortical resection. The performance of the unilateral external fixator was tested ex vivo using femoral specimens from 12-week-old Sprague-Dawley rats.

[0091] Animal surgery and DO protocol

[0092] Before surgery, each rat was anesthetized with 2–3% isoflurane (VetOne, Boise, ID, USA) during surgery on a heating pad. Cefazolin (25 mg / kg) was then injected subcutaneously into the animals. After disinfection, a 25 mm incision was made at the lateral bone site of the left femur. Four fixation pins were inserted into the bone for fixation. Another fixation pin was inserted into the first cortical bone only for bone transport, and then a frame was installed to fix the bone. A transverse femoral osteotomy was performed at the midshaft of the femur using a wire diamond saw to remove an 8 mm bone slice under sterile conditions. In addition, a 4 mm transverse osteotomy was performed at the distal femur for bone transport. An IM nail device with a length of 18 mm and a diameter of 1.2 mm was inserted into the medullary canal of the proximal and distal femoral segments, and both ends of the IM nail device were fixed with a proximal and distal fixation pin. As described above, IM nails were coated with hydrogels containing 0 μg (IMN, 16 cases), 2 μg (IMN+B2, n=16), or 6 μg (IMN+B6, n=8) of BMP-2. Animals without IM nail implantation served as blank controls (BLK, n=16). The surgical incisions were then sutured sequentially. The treatment period of the DO protocol in this study consisted of three phases: a 5-day waiting period (from the day of surgery to postoperative day (POD) 5), an 8-day active distraction period (0.5 mm / 12 h, from POD5 to POD13), and a 21-day (POD13 to POD34) or 42-day (POD13 to POD55) bone consolidation period. In this animal study, bone transport was performed in a retrograde direction through the IM nail.

[0093] Post-operative care and specimen collection

[0094] After surgery, animals were housed individually. All animals were monitored and cared for pin track infection during the study period. Polyvinylpyrrolidone iodine and ethanol were used during pin track care as necessary. Pin track infection was monitored and graded according to the Checketts-Otterburn classification. According to this classification method, pin track infection is classified into two groups: mild (grades 1-3) and severe (grades 4-6). Three rats from each group were randomly selected and subcutaneously injected with xylenol orange (30 mg / kg; Sigma-Aldrich, St. Louis, MO, USA) 13 days before termination (POD33) and calcein (10 mg / kg; Sigma-Aldrich, St. Louis, MO, USA) for in vivo labeling 3 days before termination (POD44). The immobilizer was removed 1 week before termination (POD48), and the animals' general ambulatory ability was observed on POD54. All animals were sacrificed and both femurs were harvested on POD 34 or POD 55 for qualitative and quantitative evaluation. Of note, in the (IMN+B6) group, all eight rats were sacrificed on POD 34. Bone union was primarily evaluated by gross observation of the specimens and further confirmed by micro-CT analysis as described below. Specimens were fixed in 10% buffered formalin for 48 h and transferred to 70% ethanol for storage.

[0095] Dynamic X-ray Imaging

[0096] Dynamic changes in bone defect healing were monitored weekly in all animals from POD5 to POD34 using a LAGO-X in vivo imaging system (Spectral Instrument Imaging, Tucson, AZ, USA). First, animals were anesthetized with 3% isoflurane in an induction chamber. Then, each animal was transferred to a heated imaging platform and positioned upside down before imaging. Imaging parameters were set at 40 keV and 18 s exposure time.

[0097] Micro-computed tomography (CT) examination

[0098] Microstructural changes in the stretch reconstruction and docking sites of animals were qualitatively and quantitatively evaluated using micro-CT. Briefly, all specimens were imaged using a Skyscan 1276 micro-CT (Bruker, Kontif, Belgium) at a voltage of 70 kV, current of 200 A, rotation step of 0.8°, and 360° scanning mode with a custom isotropic resolution of 20 μm isotropic voxel size. Beam hardening reduction was performed using a 0.5 mm Al filter. Projection images were reconstructed offline using the cone beam NRecon application (version 1.0.7.0, Bruker) with post-alignment and beam hardening correction for image analysis. Post-processing of the reconstructed images was analyzed using the SkyScan CAn software package (version 1.17, Bruker). Two regions of interest (ROIs), including the extension region (8 mm long) and the docking site (1.5 mm long), were analyzed separately. Cross-sectional slices of the extension region or docking site were used for bone tissue volume fraction (bone volume / total volume, BV / TV) measurements by CTAn. BW / TV of ROIs were normalized by the contralateral intact control. 3D bone structures were created from the segmented datasets of CTAn (CT Hounsfield Units (HU) threshold >10000) for visual inspection using MicroView3D image viewer (version 2.5.0; Parallax Innovations Inc., Ilderton, Canada). Nonunion was confirmed by microCT analysis defined by the absence of radiographic evidence of bone union at the reconstruction or docking site.

[0099] Mechanical testing

[0100] After micro-CT analysis, a three-point bending test was performed within 24 hours to evaluate the mechanical properties. The femurs were fracture tested using a material testing system Instron5944 Test System (Norwood, MA, USA) equipped with a 2 kN load cell. The femurs were loaded in the anterior-posterior direction with the medial and lateral spans of the blade set at 8 mm and 18 mm, respectively. The bones were tested at a speed of 0.01 mm / s, and the long axis of the femur was placed perpendicular to the blade during testing. The tensile modulus of elasticity (Young's modulus), maximum load rate, and fracture energy (energy absorption rate) were determined and analyzed with built-in software (OMAT Professional; Tinius Olsen, Inc., Horsham, PA, USA). The biomechanical properties of the new bone were expressed as a percentage of the properties of the contralateral intact bone. During mechanical testing, we terminated the compression test when the load decreased by 15% to prevent bone fracture.

[0101] histology

[0102] Immediately after mechanical testing, specimens were fixed in 10% formalin for 48 h and then transferred to 70% ethanol. Samples (n=5 / group) were decalcified in 10% EDTA solution for 5 weeks, dehydrated in ethanol, and embedded in paraffin. Thin sections (5 μm) were cut along the long axis of each femur in the sagittal plane using a microtome (RM2525; Leica, Wetzlar, Germany). Slides were stained with hematoxylin and eosin (H&E; Sigma-Aldrich) or Masson's trichrome stain (Abcam, Cambridge, MA, USA) according to standard protocols.

[0103] Histomorphometry

[0104] For histomorphometry, a protocol of paraffin embedding of mineralized bone was applied in this study. After fixation, some samples (n=3) were treated with 5.0% (w / v) potassium hydroxide aqueous solution for 96 h at room temperature on an orbital shaker. Then, the bones were washed with water and dehydrated in ethanol under a vacuum infiltration system. The treated bones were routinely embedded in paraffin-embedded blocks. 10 μm sections were cut with a RM2255 microtome (Leica, Germany) along the long axis of each femur in the sagittal plane. For histomorphometry, two sections spaced 100 μm apart were selected for measurement. Fluorescent images were taken with an all-in-one fluorescent microscope BZ-X800 (Keyence, Osaka, Japan). Five random images were taken at the docking or reconstruction sites at 10x magnification and used for measurement. The mineral accretion rate (MAR) was calculated by dividing the distance between the red and green labels by a 10-day interval.

[0105] Immunofluorescence and immunohistochemistry assays

[0106] Immunofluorescence and immunohistochemistry assays were performed using standard protocols. Samples were incubated overnight at 4°C with anti-osteocalcin (Santa Cruz, Dallas, TX, USA), anti-BMPRII (Santa Cruz, Dallas, TX, USA), or anti-α-SMA (Santa Cruz, Dallas, TX, USA) antibodies. For immunofluorescence, Alexa Fluor 488, 594, or 647-labeled secondary antibodies (Abcam, Cambridge, MA, USA) were used. For immunohistochemistry, a horseradish peroxidase-streptavidin detection system (Dako, Santa Clara, CA, USA) was used, followed by counterstaining with hematoxylin.

[0107] statistical analysis

[0108] All quantitative data were presented as mean and standard deviation (SD). After confirming normal distribution with the Kolmogorov-Smirnov test, all parameters were analyzed by ANOVA and post-hoc Tukey's HSD test. For micro-CT analysis and mechanical testing, parameters were normalized using the contralateral femur. Statistical analysis was calculated using SPSS (version 16.0; SPSS Inc., Chicago, IL, USA) and the significance level was set at P < 0.05.

Claims

1. 1. A method for producing a bioactive orthopedic implant for reconstructing a bone defect, bone deformity, or bone nonunion, comprising: (i) providing a rod and treating the surface of the rod to modify the surface properties of the rod; (ii) physically cross-linking the surfaces of the rods via charged polymer and salt ions to form a lyophilized hydrophilic hydrogel network; (iii) entrapment and loading of biological material in the hydrophilic hydrogel network; (iv) chemically cross-linking the hydrophilic hydrogel network with covalently reactive macromonomers to enhance the physical cross-linking of the hydrophilic hydrogel network to itself and to the surface of the rod.

2. 10. The method of claim 1, (v) the method further comprising the step of forming one or more coating layers on the surface of the rod of the bioactive orthopedic implant.

3. 10. The method of claim 1, the rods are interconnected porous rods; The method wherein the biological material is carried in the pores of the interconnected porous rods.

4. 10. The method of claim 1, The method, wherein the bioactive orthopedic implant is sized to fit for intramedullary implantation between the two bone segments, within the bone tunnel, or at the fracture site.

5. 1. A bioactive orthopedic implant comprising: (i) a rod having a surface that has been treated to modify its surface properties; (ii) a freeze-dried hydrophilic hydrogel network physically crosslinked to the surface of the rod via charged polymer and salt ions; (iii) a biological material entrapped in, and thereby supported by, the hydrophilic hydrogel network; and (iv) a bioactive orthopedic implant comprising the hydrophilic hydrogel network itself and a covalently bonded reactive macromonomer chemically cross-linked to the hydrophilic hydrogel network to enhance the physical cross-linking of the hydrophilic hydrogel network to the surface of the rod.

6. 6. The bioactive orthopedic implant of claim 5, the surface of the rod is coated with a covalent molecule; The bioactive orthopedic implant, wherein the covalent molecules are chemically crosslinked to the covalently reactive macromonomers to enhance adhesion of the hydrophilic hydrogel network, which is chemically and physically crosslinked to the surface of the rod, to the rod.

7. 6. The bioactive orthopedic implant of claim 5, the rods are interconnected porous rods; A bioactive orthopedic implant, wherein the biological material is carried in the pores of the interconnected porous rods.

8. 6. The bioactive orthopedic implant of claim 5, The bioactive orthopedic implant is sized to be suitable for intramedullary implantation between two bone segments, within a bone tunnel, or at a fracture site.

9. 6. The bioactive orthopedic implant of claim 5, The bioactive orthopedic implant further comprises one or more coating layers.

Images