Polycaprolactone and collagen substance compositions

Abstract

The invention provides polycaprolactone and collagen substance compositions, as well as methods of making and using them.

Description

POLYCAPROLACTONE AND COLLAGEN SUBSTANCE COMPOSITIONSCROSS-REFERENCE TO RELATED APPLICATIONS

[0001] Pursuant to 35 U.S.C. § 119(e), this application claims priority to the filing date of United States Provisional Patent Application Serial No. 63 / 546,483 filed October 30, 2023, the disclosure of each of which is herein incorporated by reference in their entirety.GOVERNMENT RIGHTS

[0002] Part of this work was performed at the Stanford Nano Shared Facilities (SNSF), supported by the National Science Foundation under award ECCS-2026822. Graded dehydration and drying was performed at the Stanford Cell Sciences Imaging Facility (CSIF). This work was supported in part by the National Science Foundation Graduate Research Fellowship (NSF GRFP) and by the National Institutes of Health under awards R01 AR074458 and R01 AR072613.BACKGROUND OF THE INVENTION

[0003] Tissue engineering is a rapidly developing field which first began with the development of artificial skin comprising collagen and glycosaminoglycans for wound healing [1][2]. The term "tissue engineering" was coined a few years later by scientists developing an ophthalmic implant made of synthetic polymers, polytetrafluoroethylene (PTFE) and polymethylmathacrylate (PMMA), which caused the eye to develop an endothelial-like membrane on the posterior surface of the prosthesis as an "engineered" cellular response to minimize light scattering from the implant material [3]. In present day, tissue engineering refers to the combination of cells, scaffolding materials, and bioactive molecules to target regenerative growth of a specific tissue. Bone tissue engineering in particular has several key challenges. Over 7 million patients are diagnosed with orthopaedic injury in the US every year, and close to 90% of trauma patients have orthopaedic injuries [4] [5]. New technological developments have enabled the treatment of large bone defects, often referred to as critical size bone defects, which do not spontaneously heal on their own [6]. However, these treatments still face issues with infection, implant loosening, and multiple surgeries. The ideal bone tissue engineering implant would be biocompatible andbioresorbable, promote the regeneration of native tissue, and degrade at a controlled rate that matches the healing rate of the native tissue.

[0004] 3D printing is a fast-emerging technology that enables scientists to precisely design and fabricate constructs to target specific applications. Among the different 3D printing methods, which include photocrosslinking (stereolithography, digital light projection) and extrusion-based approaches (molten material extrusion, syringe, inkjet), molten material extrusion (MME) has the advantage of creating large, customizable scaffolds and constructs for critical size bone defects with a relatively fast turnaround time and low costs. This makes MME an ideal approach for fabricating large bone tissue engineering constructs.

[0005] Among different biomaterial candidates for bone tissue engineering, synthetic materials, such as polymers and polyesters, have gained significant traction in the past decade. Poly lactic acid (PLA) and poly(lactic-co-glycolic acid) (PLGA), for instance, have been used commercially for drug delivery devices [7]. Among these polymers, polycaprolactone (PCL) is an FDA-approved, biocompatible polymer that was first synthesized in 1930 and gained traction in the field of tissue engineering as a bioinert polymer that can be functionalized by blending with natural biomaterials in the absence of harsh crosslinking chemicals [8]. The benefits of PCL include low immunogenicity and fewer acidic degradation byproducts compared in vivo to its counterparts, such as PLA and PLGA [9]. However, this is tied to its slow degradation in vivo, which occurs on the order of 1 to 4 years depending on the composition and crystallinity, and is too long for many applications, which target degradation time frames of 4 to 6 months [8]

[0010] . Naturally derived materials, on the other hand, typically offer high biocompatibility at the expense of fast degradation and variable mechanical properties. Gelatin is one such example and is a partially hydrolyzed, denatured form of collagen that contains arginine-glycine-aspartic acid (RGD) functional moiety sequences that improve cell attachment, differentiation, and proliferation

[0011] , Gelatin has been shown to exhibit low immunogenicity and has been characterized extensively as a biomaterial for a number of applications, including vascular tissue engineering, skin wound healing, neural cell alignment, and bone tissue repair

[0012]

[0013]

[0014]

[0011] . Although gelatin has a lower immunogenicity in vivo than collagen and is more easily handled due to its reduced temperature sensitivity, gelatin alone faces the challenge of having weak structural stability inphysiological and aqueous conditions without the addition of crosslinking and stabilizing agents, which are often cytotoxic

[0015] .

[0006] Thus, the combination of PCL and gelatin aims to create an improved biomaterial that is bioactive, robust, and degradable. PCL and gelatin have been combined by injecting or coating gelatin hydrogel onto a rigid PCL scaffold

[0016]

[0017]

[0018] or by reinforcing gelatin hydrogels with 3D printed microfibers of PCL

[0019] . However, homogeneous blends of PCL and gelatin have been primarily fabricated through electrospinning, which involves the electrification of a liquid droplet through a nozzle to form a jet that solidifies into fibers, which are then collected on a conductive plate

[0020] . Electrospun PCL and gelatin blends have been shown to improve cell growth, alignment, and attachment

[0021]

[0022]

[0015] .SUMMARY OF THE INVENTION

[0007] In a first aspect, the invention provides a homogeneous bulk PCL-Gelatin (PG) composite that contains nanosized gelatin particles that are able to withstand repeated exposure to high temperatures (>100°C) while retaining their bioactivity. This bulk PG composite material can be further processed into a 3D printing filament, which can be wound into a spool and stored until use. This method of solid PG fabrication and storage enhances the versatility and manipulability of the composite material and avoids several of the limitations of electrospinning, including the dependence of fiber and construct quality on dynamic liquid solution properties, such as viscosity, volatility, stability, and conductivity

[0023]

[0024] , Electrospun constructs are also typically limited in geometry to mats, sheets, or tubes and comprise closely packed scaffold fibers that are difficult to precisely control, which often leads to loss of interconnected pores and poor cell infiltration and migration

[0023] . Blends of PCL and gelatin also face challenges with residual solvent toxicity, which are difficult to address in situ in electrospun fibers and films without intensive post-processing steps such as washing, freeze-drying, and heat cycling, which can change the physical and chemical properties of the constructs and negatively affect their structural integrity. This solid PG fabrication methodology homogenizes the materials in the liquid phase and removes residual solvents in upstream processes, leading to high biocompatibility and tunability downstream with 3D printed constructs, both in vitro and in vivo. The creation of 3D filament also enables the design and fabrication of custom constructs, including but not limited to 3D reconstructions, porous scaffolds, meshes, and thinfilms, for a range of biomaterial and tissue engineering applications both within and outside of orthopaedics.BRIEF DESCRIPTION OF THE DRAWINGS

[0008] FIG 1 shows PCL-Gelatin material synthesis for 3D printing. PCL and gelatin are dissolved in HFIP, and 0.05% glacial acetic acid is added to reduce phase separation. The clear, homogenized mixture is then cast in a PTFE dish, evaporated of solvent to form a material sheet, and cut into pellets. PG pellets are purged of residual acetate salts by a repeated soak in cold water (4 °C) prior to freeze drying. Filament is then extruded for 3D printing.

[0009] FIG 2 shows 3D printed PCL-Gelatin scaffolds with varying porosities. From left to right: 0, 20, 40, 60, and 80 percent porosity.

[0010] FIG 3 shows (a) Instron uniaxial tensile testing setup, (b) Young’s moduli of PG groups from uniaxial tensile testing data, (c) Peak forces of PG groups from uniaxial testing data, (d) Compressive stiffnesses of PG groups from uniaxial compression data of 3D printed scaffolds of different porosities.

[0011] FIG 4 shows SEM images of bulk PG material and cross sections, (a) The hPCL surface shows a smooth, uniform matrix, (b) The PG 10 group shows a small distribution of embedded gelatin nanoparticles, (c) PG20 exhibits a higher density of embedded gelatin particles, (d) The PG30 sample shows a dense and homogeneous distribution of nano-sized gelatin particles within the PCL matrix, (e) The PGT-712 group, which contains 10% gelatin and 20% P-TCP, shows both spherical gelatin nanoparticles and irregularly shaped calcium phosphate nanoparticles, (f) PCL-TCP at an 80 / 20 ratio is shown as the control with irregular P-TCP nanoparticles. All scale bars correspond to 1 pm.

[0012] FIG 5 shows contact angle measurements on 3D printed PG disks using 4 pL water droplets. Increasing gelatin content corresponds to higher hydrophilicity.

[0013] FIG 6 shows (a) TGA setup with a material segment sample, (b) TGA of PG composite materials and controls. Curves show the thermal degradation of PCL, gelatin, and P-TCP components, with a sharp decrease in mass occurring at 250 °C and 300 °C for pure gelatin and PCL, respectively, and in the range of 350 to 400 °C for the PG composite, (c) Samples after a TGA run. PGT-712 is shown on the left, with the 20% remaining mass a white P-TCP nanopowder. The right is a representative imageof PCL, PG, and gelatin materials, which leaves a tar-like residue at the end of a ramp test.

[0014] FIG 7 shows (a) Absorption spectra of PG materials at an infrared range of 4000 cm1to 500 cm1. Peaks can be seen at 1657 cm1, and 1544 cm'1for gelatinpositive groups, and peak shortening and peak broadening observed at 1721 and 1162 cm'1, respectively, indicating the synergistic effects of gelatin in the composite materials, (b) Window of absorption spectra to identify functional group absorption frequencies (4000 cm1to 1500 cm'1). Peaks are seen at 3311 cm'1, 1657 cm'1, and 1544 cm'1, corresponding to the amine and amide groups of gelatin.

[0015] FIG 8 shows (a) Accelerated degradation of PG in alkaline solution occurs via hydrolysis, beginning with surface-initiated crack formation and followed by radial disassembly of the composite polymer network. In filaments, this results in a gummy surface and progressively thinner diameter, wherein the disassembled PG disperse into the alkaline media, forming a cloudy solution, (b) Filament prior to accelerated degradation (left) and after 6 hours in NaOH with visibly thinner diameter (right).(c) Top: SEM image of surface cracks formed after 3h of NaOH immersion at 37 °C. Bottom: High magnification of the highly porous PG surface after 3h of NaOH treatment, with cavities where gelatin nanoparticles were leached from the PG network, (d) Accelerated degradation in alkaline solution. Higher gelatin content corresponds to significantly faster degradation, whereas P-TCP content moderately increases degradation rate, (e) Degradation kinetics of composite materials in PBS simulating physiological conditions (pH 7.4) over the span of 6 weeks. The PG30 group (green square) shows a significantly faster degradation in PBS compared to the control (hPCL, blue diamond), even in the absence of enzymatic activity. AU groups are presented as mean and s.e.m. (N=5).

[0016] FIG 9 shows a primed media study to assess gelatin bioactivity after repeated exposure to high temperatures (100 °C, 3 cycles). Media was primed by immersing (a) PCL, (b) PG30, and (c) dissolved gelatin type A (uncrosslinked) in DMEM for 24h prior to the experiment. hMSCs were suspended in the primed media and seeded on ultra low adhesion well plates, and cell attachment and morphology were visually assessed after 1 day. Cell attachment can be seen in all gelatin-positive groups, indicating that the gelatin released from the PG materials retain their bioactivity, coating the bottom of the well and allowing for cell attachment and elongation. Largecell aggregates can be seen in the negative control group (a), indicating that the cells cannot sufficiently attach to the low adhesion surface. A mixture of cell attachment and aggregation is seen in the positive control group (c).

[0017] FIG 10 shows a primed DMEM media study comparing hMSC growth on tissue culture plastic well plates when indirectly exposed to PCL (top) vs. PG30 (bottom) filament through a Transwell membrane (0.4pm porosity) over 8 days. The cells exposed to the PG30 group proliferated significantly faster and reached full confluence by Day 4, peeling off of the edges of the well plate by Day 6, compared to those exposed to PCL, which reached full confluence at Day 6 and began peeling by Day 8.

[0018] FIG 11 shows (a) Experimental setup. Bone marrow-derived mesenchymal stem cells were seeded onto 3D printed disks, (b) MTS assay of human mesenchymal stem cells seeded on 3D printed PG disks and cultured in FBS -positive Dulbecco’s Modified Eagle Medium for 1 week, (c-g) SEM images of 3D printed disks seeded with hMSCs and cultured for 1 week prior to fixation, (c) hMSC attached to hPCL substrate, (d) Cells on PG10 group appear more 3-dimensional compared to the control, (e) Cells on PG20 group show extracellular vesicle formation, (f) PG30 group shows hMSCs with elongated morphology, higher filopodia count, and extracellular vesicle formation, (g) PGT-712 group shows more elongated cell morphology with dense extracellular vesicles, (h) PCL-TCP group shows more filopodia and extracellular vesicles than the PCL control, (i-h) hMSCs on hPCL substrate (i) at lower magnification shows flatter morphology compared to those on PG30 (j) and PCL-TCP (k), which appear more 3D, with the PG30 group showing more filopodic protrusions, three dimensionality, and extracellular vesicle formation. AU gelatinpositive groups have voids in the material where gelatin particles have escaped from the PG matrix. Scale bars are 10 pm for (c-h) and 100 pm for (i-k).

[0019] FIG 12 shows (a) Alkaline phosphatase staining on 3D printed disks after 1, 11, and 14 days of culture, (b) Quantification of relative change in alkaline phosphatase activity (N=3) on Day 11 and Day 14. (c) Elongated cell morphology and the appearance of cell-to-cell interactions with the PGT-712 group on Day 14. Scale bar = 40 pm.

[0020] FIG 13 shows (a) Filament implant material groups, (b) surgical implantation, and (c) overall degradation rate over the period of the 24-week study, as measured by change in postoperative vs. pre-operative volume.

[0021] FIG 14 shows histological sections of H&E stained tissue surrounding the subcutaneous filament implants after 24 weeks, (a) PCL, (b) PG 10, (c) PG20, (d) PG30, and (e) PGT-712. All scale bars represent 60 m.

[0022] FIG 15 shows live X-ray images of the rat femoral bone defect region over the span of 28 days. In the first week (Day 1 and 7), there is little calcified tissue visible. At the two- week mark, PG30 and PCL-TCP show a greater area of hard tissue formation, which continues through Day 21. At the last time point, Day 28, the amount of calcified tissue appears visually comparable among the three groups.

[0023] FIG 16 shows (a) p-CT images of the rat femoral defect at POD 28. (b) Mechanical testing of the femur explants and -CT quantification of the bone volume and tissue volume fraction. p-CT quantification shows no significant difference in the bone volume vs. tissue volume ratio. The mechanical strength of the bone, as evaluated by maximum load to failure, Young’s Modulus, and energy absorption, from the PG30 group appears to be significantly higher compared to the PCL-TCP control. The y-axis plots average percentage for each material group, normalized to the value at the initial time point (Day 1). N=4 for PCL and PG30 and N=3 for the PCL-TCP group.

[0024] FIG 17 shows histological sections of stained tissues (H&E, top, Masson’s trichrome, bottom) surrounding the rat femoral bone defect region after 28 days. OB indicates native (old) bone, and NB indicates newly formed bone. Arrows indicate preosteoblast cells and osteoids at the interface of newly formed and mature bone.

[0025] FIG 18 shows (a) Scaffold appearance, (b) Micro CT image of scaffold. 3D printed PG scaffold has approximately 3.5 mm lumen, 10 mm in diameter, and 30 mm length, (c) Image of scanning electron microscope.

[0026] FIG 19 shows a surgical view and the method of the scaffold reduction rate, (a) Surgery view at scaffold group. 3 cm bone defect was created in the femur with plate fixation, and the scaffold was inserted into the defect. The scaffold was surrounded with nylon thread along with the plate, (b) Representative image ofdecomposition ratio evaluation. Postoperative resin sections with no staining were compared to preoperative pCT image about the average of five scaffold widths.

[0027] FIG 20 shows radiographic images, (a) X-ray at 0,4,8 weeks after surgery. The black arrows point to new bone between the defects. In the PG and PG+BMP2 groups, new bone was seen inside and surrounding the scaffold, (b) Micro CT and 3D images after Microfil injection at 8weeks after surgery. New bone (white arrows) was forming as it entered the tunnel in the scaffold. Triangle arrows indicate a vessel filled with Microfil. were seen to enter the gap within and surround the scaffold. There were vessels to enter the structure of the scaffold, and vessels outside the scaffold and inside the scaffold were connected.

[0028] FIG 21 shows histology images stained with Van Gieson's picrofuchsin and SteveneTs blue staining, (a) In the Empty group, the defect was covered with soft tissue and there was disorganized new bone, (b, c) New bone was growing to enter the lumen of the scaffold and around the scaffold (yellow arrow). Within the scaffold, a vessel filled with microfil was seen. Yellow bar:5mm, red bar: 1mm.

[0029] FIG 22 shows an assessment of the amount of vessel and bone volume, (a) New bone volume, (b) Ratio of vessel inside scaffold, (c) Ratio of vessel inside scaffold, excluding scaffold volume, (d) Ratio of vessel outside scaffold, (e) Reduction rate of scaffold width. The mean reduction rates were 4.4±2.6 % in the PG group and 4.2+1.2 % in the PG+BMP-2 group, (f) The average of percentage of blood vessels in each region.DETAILED DESCRIPTION OF THE INVENTIONI. Definitions and Abbreviations

[0030] A “heat extrusion composition” can be a solid or semi-solid composition (less than 1, 2, 3, 4, or 5% solvent) which can be subjected to heat energy and mechanical energy (e.g. moving through a nozzle). The amount of heat energy can place the composition above its glass transition temperature. The amount of heat energy for polycaprolactone can be above 60°C, or above 85 °C, or above 90°C. The amount of heat energy for a polycaprolactone and collagen substance composition described herein can be above 90°C, or above 100°C, or above 105°C, or above 110°C, or above 115°C, or above 120°C, or between about 95°C and about 150°C, or between about 100°C and about 120°C.

[0031] An “additive manufacturing composition” can mean a manufacturing technology as defined in the international standard ASTM 2792-12, which mentions a process of joining materials to make objects from 3D model data, usually layer upon layer, as opposed to subtractive manufacturing methodologies, such as traditional machining. The additive manufacturing method may be selected in, but is not limited to, the list consisting of stereo lithography, mask stereolithography or mask projection stereolithography, polymer jetting, scanning laser sintering or SLS, scanning laser melting or SLM, fused deposition modeling or FDM. Additive manufacturing technologies comprise processes which create objects by juxtaposition of volume elements according to a pre-determined arrangement that can be defined in a CAD (Computer Aided Design) file. Such juxtaposition is understood as the result of sequential operations such as building a material layer on top of a previously obtained material layer and / or juxtaposing a material volume element next to a previously obtained volume element. In certain instances the additive manufacturing composition is subjected to the same temperatures as mentioned herein for the heat extrusion composition.II. Introduction

[0032] The invention provides compositions, as well as methods of making and using them.III. CompositionIll.a) Heat Extrusion Composition

[0033] In one aspect, the invention provides a heat extrusion composition comprising polycaprolactone (“PCL”) and a collagen substance. In an exemplary embodiment, the heat extrusion composition is a solid. In certain embodiments, the heat extrusion composition has not yet been subjected to heat extrusion. In certain embodiments, the heat extrusion composition has been heat extruded. Examples of heat extrusion processes include fuse deposition modeling (FDM), fused filament extrusion (FFF), molten metal deposition (MMD), and molten metal extrusion (MME). In certain instances, the heat extrusion composition is drawn through a nozzle, where it is heated (e.g. be above 90°C, or above 100°C, or above 105°C, or above 110°C, or above 115°C, or above 120°C, or between about 95°C and about 150°C, or between about 100°C and about 120°C) and is then deposited layer by layer to form a heat extrudedarticle and / or heat extruded object. The nozzle can move horizontally and a platform moves up and down vertically after each new layer is deposited.

[0034] In an exemplary embodiment, the PCL can be purchased from sources such as Sigma and Evonik. In an exemplary embodiment, the collagen substance in the heat extrusion composition is selected from the group consisting of collagen, methacrylated collagen, gelatin, and methacrylated gelatin. In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 99: 1 (w / w) to 60:40 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 99: 1 (w / w) to 80:20 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 90:10 (w / w) to 80:20 (w / w), or from 86: 14 (w / w) to 84: 16 (w / w), or from 80:20 (w / w) to 70:30 (w / w), or from 76:24 (w / w) to 74:26 (w / w), or from 95:5 (w / w) to 85: 15 (w / w), or from 85:15 (w / w) to 75:25 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 95:5 (w / w) to 65:35 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 95:5 (w / w) to 80:20 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 95:5 (w / w) to 85:15 (w / w), or from 93:7 (w / w) to 87:13 (w / w), or from 91:9 (w / w) to 89: 11 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 85: 15 (w / w) to 75:25 (w / w), or from 83:17 (w / w) to 77:23 (w / w), or from 81: 19 (w / w) to 79:21 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 75:25 (w / w) to 65:35 (w / w), or from 73:27 (w / w) to 67:33 (w / w), or from 71:29 (w / w) to 69:31 (w / w).

[0035] In an exemplary embodiment, the heat extrusion composition further comprises an additive. In an exemplary embodiment, the additive is selected from the group consisting of a bioceramic additive, an anti-pathogenic additive, and an imaging additive. In an exemplary embodiment, the additive is selected from the group consisting of beta-tricalcium-phosphate, hydroxyapatite, calcium phosphate, biphasic calcium phosphate, aluminum oxide, zirconium dioxide, titanium oxide, zinc oxide, copper oxide, and magnesium oxide. In an exemplary embodiment, the additive is beta-tricalcium-phosphate. In an exemplary embodiment, the ratio of the PCL to the collagen substance to the additive is from 70:10:20 (w / w / w) to 30:50:20 (w / w / w).

[0036] In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker derived from an aldehyde or an epoxide. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker derived from hydroxysuccinimde, maleimide, norbomene, acrylate, or methacrylate. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker catalyzed by an enzyme such as transglutaminase. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker derived from genipin and phenolic acid. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinking moiety, which is an imine or an ether.

[0037] In an exemplary embodiment, the collagen substance is present in the heat extrusion composition in particles with a diameter of between 5 nm to 500 nm, or 10 nm to 100 nm, or 10 nm to 50 nm, or 45 nm to 100 nm, or 20 nm to 90 nm, or 20 nm to 70 nm, or 40 nm to about 65 nm. In an exemplary embodiment, between 5% to 98%, or 5% to 25%, or 45% to 85%, or 65% to 90%, of the collagen substance in the heat extrusion composition are present in the particles with a diameter.

[0038] In an exemplary embodiment, the collagen substance is distributed throughout the composition. In an exemplary embodiment, between 40% to 90%, or 40% to 80% or 50% to 75% of the collagen substance is present in the interior of the composition.

[0039] In an exemplary embodiment, the heat extrusion composition is a solid. In an exemplary embodiment, the heat extrusion composition is a semi-solid. In an exemplary embodiment, the heat extrusion composition contains less than 1% (w / w), or less than 2% (w / w), or less than 3% (w / w), or less than 4% (w / w), or less than 5% (w / w) solvent. In an exemplary embodiment, the solvent is 1 ,1 ,1 ,3,3,3-hexafluoro-2- propanol (HFIP) or trifluoroethanol (TFE). In an exemplary embodiment, the heat extrusion composition contains less than 5%, or 4%, or 3%, or 2%, or 1%, or 0.8%, or 0.6%, or 0.5%, or 0.4%, or 0.3%, or 0.2%, or 0.1% organic acid salts. In an exemplary embodiment, the heat extrusion composition contains between about 0.05% to about 5% organic acid salts. In an exemplary embodiment, the heat extrusion composition contains between about 0.05% to about 0.2% organic acid salts. In an exemplaryembodiment, the organic acid is selected from the group consisting of acetic acid, malic acid, tartaric acid, lactic acid, citric acid, and ascorbic acid. In an exemplary embodiment, the organic acid is acetic acid. In an exemplary embodiment, the acetic acid salts can be sodium diacetate or sodium acetate. In an exemplary embodiment, the heat extrusion composition is a filament. In an exemplary embodiment, the heat extrusion composition is a filament, wherein the heat extrusion composition has not yet been subjected to heat extrusion, wherein the filament diameter is between about 1 mm-3 mm or 500 pm-5 mm. In an exemplary embodiment, the heat extrusion composition is a filament, wherein the heat extrusion composition has been subjected to heat extrusion, wherein the filament diameter is between 1 pm- 10 cm or 50 pm-500 pm.

[0040] In an exemplary embodiment, the heat extrusion composition is produced by a process described herein. In an exemplary embodiment, the process comprises: (a) contacting polycaprolactone, gelatin, and an organic acid in a solvent; (b) casting the product of (a); (c) subjecting the product of (b) to evaporative conditions; (d) forming the product of (c) into pellets, thereby producing the additive manufacturing composition.Ill.b) Additive Manufacturing Composition

[0041] In one aspect, the invention provides an additive manufacturing composition comprising polycaprolactone (“PCL”) and a collagen substance. In an exemplary embodiment, the additive manufacturing composition is a solid. In certain embodiments, the additive manufacturing composition has not yet been subjected to additive manufacturing. In certain embodiments, the additive manufacturing composition has been additively manufactured. Examples of additive manufacturing composition include a lattice structure, mesh, porous graft, or overhanging structure. In certain instances, the additive manufacturing composition is drawn through a nozzle, where it is heated between about 90°C to 120°C and is then deposited layer by layer to form an additive manufactured article and / or additive manufactured object. The nozzle can move horizontally and a platform moves up and down vertically after each new' layer is deposited .

[0042] In an exemplary embodiment, the PCL can be purchased from sources such as Sigma and Evonik. In an exemplary embodiment, the collagen substance in the additive manufacturing composition is selected from the group consisting of collagen,methacrylated collagen, gelatin, and methacrylated gelatin. In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 99: 1 (w / w) to 60:40 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 99: 1 (w / w) to 80:20 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 90:10 (w / w) to 80:20 (w / w), or from 86: 14 (w / w) to 84: 16 (w / w), or from 80:20 (w / w) to 70:30 (w / w), or from 76:24 (w / w) to 74:26 (w / w), or from 95:5 (w / w) to 85: 15 (w / w), or from 85:15 (w / w) to 75:25 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 95:5 (w / w) to 65:35 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 95:5 (w / w) to 80:20 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 93:7 (w / w) to 87:13 (w / w), or from 91:9 (w / w) to 89:11 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 83:17 (w / w) to 77:23 (w / w), or from 81:19 (w / w) to 79:21 (w / w). In an exemplary embodiment, the ratio of the PCL to the collagen substance is from 73:27 (w / w) to 67:33 (w / w), or from 71:29 (w / w) to 69:31 (w / w).

[0043] In an exemplary embodiment, the invention further comprises an additive. In an exemplary embodiment, the additive is selected from the group consisting of a bioceramic additive, an anti-pathogenic additive, and an imaging additive. In an exemplary embodiment, the additive is selected from the group consisting of beta- tricalcium-phosphate, hydroxyapatite, calcium phosphate, biphasic calcium phosphate, aluminum oxide, zirconium dioxide, titanium oxide, zinc oxide, copper oxide, and magnesium oxide. In an exemplary embodiment, the additive is beta- tricalcium-phosphate. In an exemplary embodiment, the ratio of the PCL to the collagen substance to the additive is from 70:10:20 (w / w / w) to 30:50:20 (w / w / w).

[0044] In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker derived from an aldehyde or an epoxide. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker derived from hydroxysuccinimde, maleimide, norbomene, acrylate, or methacrylate. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker catalyzed by an enzyme such astransglutaminase. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinker derived from genipin and phenolic acid. In an exemplary embodiment, the collagen substance is not covalently attached to the polycaprolactone through a crosslinking moiety, which is an imine or an ether.

[0045] In an exemplary embodiment, the collagen substance is present in the additive manufacturing composition in particles with a diameter of between 5 nm to 500 nm, or 10 nm to 100 nm, or 10 nm to 50 nm, or 45 nm to 100 nm, or 20 nm to 90 nm, or 20 nm to 70 nm, or 40 nm to about 65 nm. In an exemplary embodiment, between 5% to 98%, or 5% to 25%, or 45% to 85%, or 65% to 90%, of the collagen substance in the additive manufacturing composition are present in the particles with a diameter.

[0046] In an exemplary embodiment, the collagen substance is distributed throughout the composition. In an exemplary embodiment, between 40% to 90%, or 40% to 80% or 50% to 75% of the collagen substance is present in the interior of the composition.

[0047] In an exemplary embodiment, the additive manufacturing composition is a solid. In an exemplary embodiment, the additive manufacturing composition is a semi-solid. In an exemplary embodiment, the additive manufacturing composition contains less than 1% (w / w), or less than 2% (w / w), or less than 3% (w / w), or less than 4% (w / w), or less than 5% (w / w) solvent. In an exemplary embodiment, the solvent is l,l,l,3,3,3-hexafluoro-2-propanol (HFIP) or trifluoroethanol (TFE). In an exemplary embodiment, the additive manufacturing composition contains less than 5%, or 4%, or 3%, or 2%, or 1%, or 0.8%, or 0.6%, or 0.5%, or 0.4%, or 0.3%, or 0.2%, or 0.1% organic acid salts. In an exemplary embodiment, the additive manufacturing composition contains between about 0.05% to about 5% organic acid salts. In an exemplary embodiment, the additive manufacturing composition contains between about 0.05% to about 0.2% organic acid salts. In an exemplary embodiment, the organic acid is selected from the group consisting of acetic acid, malic acid, tartaric acid, lactic acid, citric acid, and ascorbic acid. In an exemplary embodiment, the organic acid is acetic acid. In an exemplary embodiment, the acetic acid salts can be sodium diacetate or sodium acetate. In an exemplary embodiment, the additive manufacturing composition is a filament. In an exemplary embodiment, the additive manufacturing composition is a filament, wherein the additive manufacturing composition has not yet been subjected to heat extrusion, wherein the filamentdiameter is between about 1 mm- 3 mm or 500 |im-5 mm. In an exemplary embodiment, the additive manufacturing composition is a filament, wherein the additive manufacturing composition has been subjected to heat extrusion, wherein the filament diameter is between 1 pm- 10 cm or 50 pm-500 pm.

[0048] In an exemplary embodiment, the additive manufacturing composition is produced by a process described herein. In an exemplary embodiment, the process comprises: (a) contacting polycaprolactone, collagen substance, and an organic acid in a solvent; (b) casting the product of (a); (c) subjecting the product of (b) to evaporative conditions; (d) forming the product of (c) into pellets, thereby producing the additive manufacturing composition. like) Methods of making

[0049] In an exemplary embodiment, the invention provides a method of making a composition described herein, the method comprising: (a) contacting polycaprolactone, collagen substance, and an organic acid in a solvent; (b) casting the product of (a); (c) subjecting the product of (b) to evaporative conditions; (d) forming the product of (c) into pellets; (e) removal of organic acid salts from the product of (d); thereby making the composition.

[0050] In an exemplary embodiment, the contacting comprises: (al) dissolving the polycaprolactone and the collagen substance in the solvent; (a2) adding the organic acid to the product of (al). In an exemplary embodiment, (al) is conducted from between about 1 hr to about 48 hrs, about 2 hrs to about 36 hrs, about 2 hrs to about 24 hrs, about 6 hrs to about 18 hrs, or about 12 hrs to about 24 hrs. In an exemplary embodiment, (al) is conducted at a temperature of from about 24°C to about 50°C, or from about 33°C to about 39°C, or about 36°C to about 38°C. In an exemplary embodiment, (al) is stirred. In an exemplary embodiment, (a2) is conducted from between about 1 second to about 1 hr, or about 1 second to about 30 mintues, or about 1 sec to about 1 min, or about 5 sec to about 1 min. In an exemplary embodiment, (a2) is conducted at a temperature of from about 24°C to about 50°C, or from about 33 °C to about 39°C, or about 36°C to about 38°C. In an exemplary embodiment, (a2) is stirred. In an exemplary embodiment, (al) and (a2) are together conducted from between about 1 hr to about 48 hrs, or about 2 hrs to about 36 hrs, or about 2 hrs to about 24 hrs, or about 6 hrs to about 18 hrs, or about 12 hrs to about 24 hrs. In an exemplary embodiment, the solvent is l,l,l,3,3,3-hexafluoro-2-propanol (HFIP) ortrifluoroethanol (TFE). In an exemplary embodiment, the organic acid is selected from the group consisting of acetic acid, malic acid, tartaric acid, lactic acid, citric acid, and ascorbic acid. In an exemplary embodiment, the organic acid is acetic acid. In an exemplary embodiment, the (a) further comprises contacting the caprolactone and the collagen substance with the additive, and then further contacting with the organic acid. In an exemplary embodiment, the additive is as described herein. In an exemplary embodiment, the casting of the (b) comprises placing the product of step (a) into a container. In an exemplary embodiment, the container is a non-stick container such as a container comprising a fluoropolymer such as polytetrafluoroethylene (PTFE). In an exemplary embodiment, the (b) is conducted from between about 1 hr to about 24 hrs, or about 2 hrs to about 16 hrs, or about 4 hrs to about 12 hrs, or about 6 hrs to about 10 hrs, or about 7 hrs to about 9 hrs. In an exemplary embodiment, (al) is conducted at a temperature of from about 24°C to about 90°C, from about 37°C to about 80°C, about 55°C to about 80°C, or about 65 °C to about 75°C. In an exemplary embodiment, the (b) is stirred. In an exemplary embodiment, the (b) is conducted in a low-form polytetrafluoroethylene evaporation dish at 70°C for 8 hours under stirring. In an exemplary embodiment, the average thickness of the composition at the conclusion of the (b) is between about 0.5 mm and about 10 mm, about 1 mm or about 6 mm, about 1.5 mm to about 4 mm, or about 2 mm to about 3 mm. In an exemplary embodiment, the solvent comprises less than 1%, 0.8%, 0.5%, 0.25%, or less than 0.1% of the product of (c). In an exemplary embodiment, the step (c) is conducted from between about 1 hr to about 96 hrs, or about 12 hrs to about 84 hrs, or about 18 hrs to about 72 hrs, or about 24 hrs to about 60. In an exemplary embodiment, step (c) is conducted in a desiccator or nearvacuum conditions. In an exemplary embodiment, step (c) is conducted at about 25 °C, or from about 24°C to about 90°C, from about 37°C to about 80°C, about 55°C to about 80°C, or about 65 °C to about 75 °C. In an exemplary embodiment, the average thickness of the composition at the conclusion of the (c) is between about 0.5 mm and about 10 mm, about 1 mm or about 6 mm, about 1.5 mm to about 4 mm, or about 2 mm to about 3 mm. In an exemplary embodiment, in (d) the product of step (c) is divided into pellets of approximately 2-4 mm in their dimensions. In an exemplary embodiment, in (d) the product of step (c) is cut with blades (e.g. stainless steel blades) first laterally into 2.5mm-3.5mm strips and then orthogonally to form 2.5mm-3.5mm diameter pellets. In an exemplary embodiment, step (d) furthercomprises, after the pellet formation, subjecting the pellets to evaporative conditions, such as those described herein. In an exemplary embodiment, for the (e), the product of step (d) is subjected to submerging or washing with water. In an exemplary embodiment, the temperature of the water is between about 1°C and about 10°C, about 1°C and about 7°C, about 2°C and about 7°C, or about 3°C and about 5°C. In an exemplary embodiment, the water pH is neutral, or between about 6.8 and 7.2. In an exemplary embodiment, for the (e), the product of (d) is subjected to deep freezing conditions (such as at a temperature of between about -90°C and about -50°C). In an exemplary embodiment, for the (e), the product of (d) is subjected to lyophilizing conditions. In an exemplary embodiment, for the (e), the product of (d) is subjected to submerging or washing with water with a temperature of between about 1 °C and about 7°C, then deep freezing conditions, and then lyophilizing conditions. In an exemplary embodiment, lyophilizing conditions can be conducted for several hours to several days (e.g., 8 hours to 72 hours), and / or in an airtight container with a vacuum pressure between 0.033 to 0.133 mbar.IV. Heat extruded article

[0051] In an exemplary embodiment, the invention provides a heat extruded article. In an exemplary embodiment, the heat extruded article comprises two or more layers of the heat extrusion composition described herein. In an exemplary embodiment, the heat extruded article is formed by a process described herein. In an exemplary embodiment, the heat extruded article is formed by molten material extrusion. In an exemplary embodiment, the heat extruded article has pores, and the porosity is controllable by the process described herein.

[0052] In an exemplary embodiment, the heat extruded article further comprises a signaling factor. In an exemplary embodiment, the signaling factor promotes angiogenesis and / or osteogenesis. In an exemplary embodiment, the signaling factor is a bone morphogenetic factor, such as bone morphogenetic factor 2 (BMP-2), recombinant BMP-2 (rBMP-2), human BMP-2 (hBMP-2) or recombinant human BMP-2 (rhBMP-2). In an exemplary embodiment, the heat extruded article has a ratio of the PCL to the collagen substance (such as gelatin) from 75:25 (w / w) to 65:35 (w / w), 73:27 (w / w) to 67:33 (w / w), or from 71:29 (w / w) to 69:31 (w / w), and further comprises a signalling factor, such as BMP-2 described herein. In an exemplary embodiment, the heat extruded article has a ratio of the PCL to the collagen substance(such as gelatin) from 83:17 (w / w) to 77:23 (w / w), or from 81:19 (w / w) to 79:21 (w / w), and further comprises a signalling factor, such as BMP-2 described herein. In an exemplary embodiment, the heat extruded article has a ratio of the PCL to the collagen substance (such as gelatin) from 95:5 (w / w) to 85:15 (w / w), or from 93:7 (w / w) to 87:13 (w / w), or from 91:9 (w / w) to 89:11 (w / w), and further comprises a signalling factor, such as BMP-2 described herein.

[0053] The signaling factor comprising heat extruded article can be made by (a) contacting the heat extruded article with a fibrinogen solution and then (b) contacting the product of step (a) with the signaling factor. In an exemplary embodiment, the contacting of step (b) is by injection of the signaling factor onto and / or into the heat extruded article. In an exemplary embodiment, the making of the signaling factor comprising heat extruded article further comprises, between the step (a) and the step (b), (a’), contacting the heat extruded article with a thrombin solution.V. Heat extruded object

[0054] In an exemplary embodiment, the invention provides a heat extruded object. In an exemplary embodiment, the heat extruded object is produced by a process comprising heating and extruding the heat extrusion composition described herein through a hot-melt extrusion nozzle to form an extrudate; and depositing the extrudate such that multiple layers are controllably deposited and fused forming the heat extruded object. In an exemplary embodiment, the extruding is by molten material extrusion. In an exemplary embodiment, the heat extruded object has pores, and the porosity is controllable by the process described herein.

[0055] In an exemplary embodiment, the heat extruded object further comprises a signaling factor. In an exemplary embodiment, the signaling factor promotes angiogenesis and / or osteogenesis. In an exemplary embodiment, the signaling factor is a bone morphogenetic factor, such as bone morphogenetic factor 2 (BMP-2), recombinant BMP-2 (rBMP-2), human BMP-2 (hBMP-2) or recombinant human BMP-2 (rhBMP-2). In an exemplary embodiment, the heat extruded object has a ratio of the PCL to the collagen substance (such as gelatin) from 75:25 (w / w) to 65:35 (w / w), 73:27 (w / w) to 67:33 (w / w), or from 71:29 (w / w) to 69:31 (w / w), and further comprises a signalling factor, such as BMP-2 described herein. In an exemplary embodiment, the heat extruded object has a ratio of the PCL to the collagen substance (such as gelatin) from 83:17 (w / w) to 77:23 (w / w), or from 81:19 (w / w) to 79:21(w / w), and further comprises a signalling factor, such as BMP-2 described herein. In an exemplary embodiment, the heat extruded object has a ratio of the PCL to the collagen substance (such as gelatin) from 95:5 (w / w) to 85:15 (w / w), or from 93:7 (w / w) to 87:13 (w / w), or from 91:9 (w / w) to 89:11 (w / w), and further comprises a signalling factor, such as BMP-2 described herein.

[0056] The signaling factor comprising heat extruded object can be made by (a) contacting the heat extruded object with a fibrinogen solution and then (b) contacting the product of step (a) with the signaling factor. In an exemplary embodiment, the contacting of step (b) is by injection of the signaling factor onto and / or into the heat extruded object. In an exemplary embodiment, the making of the signaling factor comprising heat extruded object further comprises, between the step (a) and the step (b), (a’), contacting the heat extruded object with a thrombin solution.VI. Additive manufactured article

[0057] In an exemplary embodiment, the invention provides an additive manufactured article. In an exemplary embodiment, the additive manufactured article comprises two or more layers of the additive manufactured composition described herein. In an exemplary embodiment, the additive manufactured article is formed by a process described herein. In an exemplary embodiment, the additive manufactured article is formed by molten material extrusion. In an exemplary embodiment, the additive manufactured article has pores, and the porosity is controllable by the process described herein.

[0058] In an exemplary embodiment, the additive manufactured article further comprises a signaling factor. In an exemplary embodiment, the signaling factor promotes angiogenesis and / or osteogenesis. In an exemplary embodiment, the signaling factor is a bone morphogenetic factor, such as bone morphogenetic factor 2 (BMP-2), recombinant BMP-2 (rBMP-2), human BMP-2 (hBMP-2) or recombinant human BMP-2 (rhBMP-2). In an exemplary embodiment, the additive manufactured article has a ratio of the PCL to the collagen substance (such as gelatin) from 75:25 (w / w) to 65:35 (w / w), 73:27 (w / w) to 67:33 (w / w), or from 71:29 (w / w) to 69:31 (w / w), and further comprises a signalling factor, such as BMP-2 described herein. In an exemplary embodiment, the additive manufactured article has a ratio of the PCL to the collagen substance (such as gelatin) from 83:17 (w / w) to 77:23 (w / w), or from 81: 19 (w / w) to 79:21 (w / w), and further comprises a signalling factor, such as BMP-2described herein. In an exemplary embodiment, the additive manufactured article has a ratio of the PCL to the collagen substance (such as gelatin) from 95:5 (w / w) to 85: 15 (w / w), or from 93:7 (w / w) to 87:13 (w / w), or from 91:9 (w / w) to 89: 11 (w / w), and further comprises a signalling factor, such as BMP-2 described herein.

[0059] The signaling factor comprising additive manufactured article can be made by(a) contacting the additive manufactured article with a fibrinogen solution and then(b) contacting the product of step (a) with the signaling factor. In an exemplary embodiment, the contacting of step (b) is by injection of the signaling factor onto and / or into the additive manufactured article. In an exemplary embodiment, the making of the signaling factor comprising additive manufactured article further comprises, between the step (a) and the step (b), (a’), contacting the additive manufactured article with a thrombin solution.VII. Additive manufactured object

[0060] In an exemplary embodiment, the invention provides an additive manufactured object. In an exemplary embodiment, the additive manufactured object is produced by a process comprising heating and extruding the additive manufacturing composition described herein through a hot-melt extrusion nozzle to form an extrudate; and depositing the extrudate such that multiple layers are controllably deposited and fused forming the additive manufactured object. In an exemplary embodiment, the extruding is by molten material extrusion. In an exemplary embodiment, the additive manufactured object has pores, and the porosity is controllable by the process described herein.

[0061] In an exemplary embodiment, the additive manufactured object further comprises a signaling factor. In an exemplary embodiment, the signaling factor promotes angiogenesis and / or osteogenesis. In an exemplary embodiment, the signaling factor is a bone morphogenetic factor, such as bone morphogenetic factor 2 (BMP-2), recombinant BMP-2 (rBMP-2), human BMP-2 (hBMP-2) or recombinant human BMP-2 (rhBMP-2). In an exemplary embodiment, the additive manufactured object has a ratio of the PCL to the collagen substance (such as gelatin) from 75:25 (w / w) to 65:35 (w / w), 73:27 (w / w) to 67:33 (w / w), or from 71:29 (w / w) to 69:31 (w / w), and further comprises a signalling factor, such as BMP-2 described herein. In an exemplary embodiment, the additive manufactured article has a ratio of the PCL to the collagen substance (such as gelatin) from 83:17 (w / w) to 77:23 (w / w), or from81: 19 (w / w) to 79:21 (w / w), and further comprises a signalling factor, such as BMP-2 described herein. In an exemplary embodiment, the additive manufactured object has a ratio of the PCL to the collagen substance (such as gelatin) from 95:5 (w / w) to 85: 15 (w / w), or from 93:7 (w / w) to 87:13 (w / w), or from 91:9 (w / w) to 89:11 (w / w), and further comprises a signalling factor, such as BMP-2 described herein.

[0062] The signaling factor comprising additive manufactured object can be made by(а) contacting the additive manufactured object with a fibrinogen solution and then (b) contacting the product of step (a) with the signaling factor. In an exemplary embodiment, the contacting of step (b) is by injection of the signaling factor onto and / or into the additive manufactured object. In an exemplary embodiment, the making of the signaling factor comprising additive manufactured object further comprises, between the step (a) and the step (b), (a’), contacting the additive manufactured object with a thrombin solution.Bibliography:[1] IV Yannas et al. “Design of an artificial skin. II. Control of chemical composition”. In: Journal of biomedical materials research 14.2 (1980), pp. 107-132.[2] Lawrence J Bonassar and Charles A Vacanti. “Tissue engineering: the first decade and beyond”. In: Journal of Cellular Biochemistry 72.S30-31 (1998), pp. 297-303.[3] Traian V Chirila and Damien G Harkin. “An introduction to ophthalmic biomaterials and their role in tissue engineering and regenerative medicine”. In: Biomaterials and regenerative medicine in ophthalmology. Elsevier, 2016, pp. 1-14.[4] Molly P Jarman et al. “The national burden of orthopedic injury: cross- sectional estimates for trauma system planning and optimization”. In: Journal of surgical research 249 (2020), pp. 197-204.[5] Melanie Berube et al. “Low-value injury care in the adult orthopaedic trauma population: a protocol for a rapid review”. In: BMJ open 10.3 (2020), e033453.[б] Erika Roddy et al. “Treatment of critical-sized bone defects: clinical and tissue engineering perspectives”. In: European Journal of Orthopaedic Surgery & Traumatology 28 (2018), pp. 351- 362.[7] Rajeev A Jain. “The manufacturing techniques of various drug loaded biodegradable poly (lactide-co-glycolide)(PLGA) devices”. In: Biomaterials 21.23 (2000), pp. 2475-2490.[8] Ohan S Manoukian. Encyclopedia of biomedical engineering. Elsevier, 2019.[9] Sumit Murab et al. “Advances in additive manufacturing of polycaprolactone based scaffolds for bone regeneration”. In: Journal of Materials Chemistry B (2023).

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[0063] Notwithstanding the appended claims, the present disclosure is also defined by the following clauses:

[0064] Clause 1. A heat extrusion composition comprising polycaprolactone and a collagen substance, wherein the heat extrusion composition is a solid.

[0065] Clause 2. The heat extrusion composition of clause 1, wherein the collagen substance is distributed throughout the composition.

[0066] Clause 3. The heat extrusion composition of a preceding clause, wherein the collagen substance is selected from the group consisting of collagen, methacrylated collagen, gelatin, and methacrylated gelatin.

[0067] Clause 4. The heat extrusion composition of a preceding clause, wherein the collagen substance is not covalently attached to the polycaprolactone through a crosslinking moiety selected from an imine or an ether.

[0068] Clause 5. The heat extrusion composition of a preceding clause, wherein the collagen substance is present in nanosized particles.

[0069] Clause 6. The heat extrusion composition of a preceding clause, wherein solvent comprises less than 1% of the heat extrusion composition.

[0070] Clause 7. The heat extrusion composition of a preceding clause, further comprising an additive.

[0063] Clause 8. The heat extrusion composition of a preceding clause, wherein the additive is selected from the group consisting of a bioceramic additive, an anti- pathogenic additive, and an imaging additive.

[0071] Clause 9. The heat extrusion composition of a preceding clause, wherein the additive is selected from the group consisting of beta-tricalcium-phosphate, hydroxyapatite, calcium phosphate, biphasic calcium phosphate, aluminum oxide, zirconium dioxide, titanium oxide, zinc oxide, copper oxide, and magnesium oxide.

[0072] Clause 10. The heat extrusion composition of a preceding clause, wherein the additive is beta-tricalcium-phosphate.

[0073] Clause 11. The heat extrusion composition of a preceding clause, wherein the composition is a filament.

[0074] Clause 12. The heat extrusion composition of a preceding clause, produced by a process comprising: (a) contacting polycaprolactone, collagen substance, and an organic acid in a solvent; (b) casting the product of (a); (c) subjecting the product of (b) to evaporative conditions; (d) forming the product of (c) into pellets, thereby producing the additive manufacturing composition.

[0075] Clause 13. A method of making the heat extrusion composition of a preceding clause, said method comprising: (a) contacting polycaprolactone, collagen substance, and an organic acid in a solvent; (b) casting the product of (a); (c) subjecting the product of (b) to evaporative conditions; (d) forming the product of (c) into pellets, thereby making the heat extrusion composition.

[0063] Clause 14. The method of clause 13, wherein the contacting comprises: (al) dissolving polycaprolactone and collagen substance in the solvent; (a2) adding the organic acid to the product of (al).

[0076] Clause 15. The method of clause 13 or 14, wherein the solvent is 1, 1,1, 3,3,3- hexalluoro-2-propanol (HFIP), and the organic acid is acetic acid.

[0077] Clause 16. The method of clause 13, 14, or 15, wherein the ratio of polycarolactone to collagen substance is between 99: 1 and 80:20.

[0078] Clause 17. The method of clause 13, 14, 15, or 16, wherein the (a) further comprises contacting the caprolactone, collagen substance, and organic acid with an additive.

[0079] Clause 18. The method of clause 13, 14, 15, 16, or 17, wherein the additive is beta-tricalcium-phosphate.

[0080] Clause 19. The method of clause 13, 14, 15, 16, 17, or 18, wherein the (b) is cast into a low-form polytetrafluoroethylene evaporation dish at 70 °C for 8 hours under stirring.

[0081] Clause 20. The method of clause 13, 14, 15, 16, 17, 18, or 19, wherein the solvent comprises less than 1% of the product of (c).

[0082] Clause 21. The method of clause 13, 14, 15, 16, 17, 18, 19, or 20, wherein the (d) is cutting the material sheet with stainless steel blades first laterally into 3mm strips and then orthogonally to form 3mm diameter rectangular pellets.

[0083] Clause 22. A heat extruded article, comprising two or more layers of the heat extrusion composition of a preceding clause.

[0084] Clause 23. The heat extruded article of clause 22, wherein the heat extruded article is formed by molten material extrusion.

[0085] Clause 24. A method of forming a heat extruded object comprising: heating and extruding the heat extrusion composition of claim 1 through a hot-melt extrusion nozzle for form an extrudate; and depositing the extrudate such that multiple layers are controllably deposited and fused forming the additive manufactured object.

[0086] Clause 25. The method of clause 24, wherein the extruding is by molten material extrusion.

[0087] The invention is further illustrated by the Examples that follow. The Examples are not intended to define or limit the scope of the invention.EXAMPLES

[0088] The following Examples illustrate the synthesis of representative compounds used in the invention and the following Reference Examples illustrate the synthesis of intermediates in their preparation. These examples are not intended, nor are they to be construed, as limiting the scope of the invention. It will be clear that the invention may be practiced otherwise than as particularly described herein. Numerous modifications and variations of the invention are possible in view of the teachings herein and, therefore, are within the scope of the invention.EXAMPLE 1Materials and Methods

[0089] A scheme of PCL-Gelatin material synthesis for 3D printing is provided in FIG. 1.

[0090] Polycaprolactone (80,000 Mn, Sigma Aldrich, USA) and gelatin type A from porcine skin (Sigma Aldrich, USA) were weighed at designated ratios and dissolved in l,l,l,3,3,3-Hexafluoro-2-propanol (HFIP) at a 5% ratio (w / v) (99% purity, Synquest Laboratories, USA). The mixture was left in an incubator shaker for 24h at 37°C. Glacial acetic acid (Sigma Aldrich) at 0.05% (v / v) was added the next day to lower the isoelectric point of the solution and minimize phase separation between materialsbefore casting

[0025] . Beta tricalcium phosphate (P-TCP) powder with an average particle size of 100 nm (Berkeley Advanced Biomaterials, USA) was then added to select materials for bone-targeted applications and vigorously mixed for Ih before casting and periodically mixing with a polytetrafluoroethylene (PTFE) rod. The homogenized solutions were cast into low-form PTFE evaporation dishes (Fisher Scientific, USA) and left in a chemical fume hood for 3 days to evaporate the solvent. Dried PG material sheets were removed from the PTFE dishes and cut into pellets (3mm diameter) and left in the fume hood for at least 24h to fully evaporate residual solvent. Residual acetate salt was then removed by washing the pellets in cold water (4°C) until pH neutral, deep freezing at -80°C, and lyophilizing.PCL-TCP

[0091] A single batch of PCL-TCP (80 / 20 mass ratio) constitutes 40g of material. To synthesize a single sheet of PCL-TCP, polycaprolactone (PCL, Mn= 80,000 kDa, Sigma Aldrich, USA) was first dissolved in 500 mL of N,N-Dimethylformamide (DMF, Certified ACS, Fisher Scientific, USA) overnight in an incubating shaker at 60 °C. After 8h, -tricalcium phosphate (P-TCP) nanocrystals (100 nm, Berkeley Advanced Materials, USA) were slowly added to the solution and allowed to mix for 1 hour to create a homogeneous mixture prior to precipitation in a cold water bath. The precipitate, a solid mass, was then manually stretched to form a thin sheet of PCL-TCP composite material and allowed to dry in a fume hood for 48 hours to evaporate any residual water and solvent. The dried material sheet was then cut into pellets (3mm diameter) and left in the fume hood for at least 24h to fully evaporate residual DMF solvent.PCL

[0092] A single batch of PCL constitutes 40g of material. To synthesize a single sheet of PCL, 40g of polycaprolactone (80,000 Mn, Sigma Aldrich, USA) was measured out and dissolved in 800 mL of l,l,l,3,3,3-Hexafluoro-2-propanol (HFIP) at a 5% ratio (w / v). Glacial acetic acid (Sigma Aldrich) at 0.05% (v / v) was added after the solutes completely dissolved to form a translucent solution. The solution was then mixed for 1 hour in an incubating shaker at 37 °C before casting into low-form polytetrafluoroethylene (PTFE) evaporation dishes (400 mL capacity, Fisher Scientific) and stirred with a magnetic stir bar on a hot plate at 70 °C for 8 hours. The PTFE dish was removed from heat, and HFIP solvent was allowed to evaporate for 3days in a chemical fume hood. After 3 days, the dried material sheet was manually removed from the PTFE dishes and cut into pellets (3mm diameter) and left in the fume hood for at least 24h to fully evaporate residual solvent. Residual acetate salt was then removed by washing the pellets in cold water (4 °C) in the fridge until pH neutral (7.0), deep freezing at -80°C, and lyophilizing.PG10

[0093] A single batch of PCL constitutes 40g of material. To synthesize a single sheet of PCL, 36g of polycaprolactone (80,000 Mn, Sigma Aldrich, USA) and 4 g of gelatin type A (Sigma Aldrich) were measured out and dissolved in 800 mL of 1, 1,1, 3,3,3- Hexafluoro-2-propanol (HFIP). Glacial acetic acid (Sigma Aldrich) at 0.05% (v / v) was added the next day. The solution was then mixed for 1 hour in an incubating shaker at 37 °C before casting into low-form polytetrafluoroethylene (PTFE) evaporation dishes (400 mL capacity, Fisher Scientific) and stirred with a magnetic stir bar on a hot plate at 70 °C overnight. The PTFE dish was removed from heat, and HFIP solvent was allowed to evaporate for 3 days in a chemical fume hood. After 3 days, the dried material sheet was manually removed from the PTFE dishes and cut into pellets (3mm diameter) and left in the fume hood for at least 24h to fully evaporate residual solvent. Residual acetate salt was then removed by washing the pellets in cold water (4°C) in the fridge until pH neutral (7.0), deep freezing at -80°C, and lyophilizing.PG20

[0094] A single batch of PCL constitutes 40g of material. To synthesize a single sheet of PCL, 32 g of polycaprolactone (80,000 Mn, Sigma Aldrich, USA) and 8 g of gelatin type A (Sigma Aldrich) were measured out and dissolved in 800 mL of l,l,l,3,3,3-Hexafluoro-2-propanol (HFIP) at a 5% ratio (w / v). Glacial acetic acid (Sigma Aldrich) at 0.05% (v / v) was added the next day. The solution was then mixed for 1 hour in an incubating shaker at 37 °C before casting into low form polytetrafluoroethylene (PTFE) evaporation dishes (400 mL capacity, Fisher Scientific) and stirred with a magnetic stir bar on a hot plate at 70 °C overnight. The PTFE dish was removed from heat, and HFIP solvent was allowed to evaporate for 3 days in a chemical fume hood. After 3 days, the dried material sheet was manually removed from the PTFE dishes and cut into pellets (3mm diameter) and left in the fume hood for at least 24h to fully evaporate residual solvent. Residual acetate saltwas then removed by washing the pellets in cold water (4 °C) in the fridge until pH neutral (7.0), deep freezing at -80°C, and lyophilizing.PG30

[0095] A single batch of PCL constitutes 40g of material. To synthesize a single sheet of PCL, 28 g of polycaprolactone (80,000 Mn, Sigma Aldrich, USA) and 12 g of gelatin type A (Sigma Aldrich) were measured out and dissolved in 800 mL of l,l,l,3,3,3-Hexafluoro-2-propanol (HFIP) at a 5% ratio (w / v). Glacial acetic acid (Sigma Aldrich) at 0.05% (v / v) was added the next day. The solution was then mixed for 1 hour in an incubating shaker at 37 °C before casting into low form PTFE evaporation dishes (400 mL capacity, Fisher Scientific) and stirred with a magnetic stir bar on a hot plate at 70 °C overnight. The PTFE dish was removed from heat, and HFIP solvent was allowed to evaporate for 3 days in a chemical fume hood. After 3 days, the dried material sheet was manually removed from the PTFE dishes and cut into pellets (3mm diameter) and left in the fume hood for at least 24h to fully evaporate residual solvent. Residual acetate salt was then removed by washing the pellets in cold water (4°C) in the fridge until pH neutral (7.0), deep freezing at -80°C, and lyophilizing.PGT-712

[0096] A single batch of PGT-712 constitutes 40g of material. To synthesize a single sheet, 28 g of polycaprolactone (80,000 Mn, Sigma Aldrich, USA), 4 g of gelatin type A were measured out and dissolved in 800 mL of l,l,l,3,3,3-Hexafluoro-2-propanol (HFIP) overnight in an incubating shaker at 37 °C. Glacial acetic acid (Sigma Aldrich) at 0.05% (v / v) was added the next day and allowed to homogenize for 1 hour. 8 g of P-tricalcium phosphate (P-TCP) powder with an average particle size of 100 nm (Berkeley Advanced Biomaterials, USA) was measured out and added to the clear, homogenized solution. The PCL-Gelatin-TCP solution was then mixed vigorously mixed for Ih in the incubating shaker to homogenously distribute the TCP nanoparticles. The solution was cast into low-form polytetrafluoroethylene (PTFE) dishes, placed onto a hot plate set at 70 °C, and stirred with a magnetic stir bar until evaporation of the HFIP solvent resulted in a viscous solution (approximately after 4 hours). The magnetic stir bar was then manually removed, and the viscous mixture was periodically mixed (every 30 minutes) with polytetrafluoroethylene (PTFE) rod to prevent TCP nanoparticle sedimentation onto the bottom of the container. After 8hours, the PTFE dish was removed from heat, and any remaining HFIP solvent was allowed to evaporate for 3 days in a chemical fume hood. After 3 days, the dried material sheet was manually removed from the PTFE dishes and cut into pellets (3mm diameter) and left in the fume hood for at least 24h to fully evaporate any residual solvent. Residual acetate salt was then removed by washing the pellets in cold water (4°C) in the fridge until pH neutral (7.0), deep freezing at -80°C, and lyophilizing.Filament Fabrication and Processing

[0097] A filament extruder with a stainless steel barrel and custom-sized screw designed to accommodate <5mm particles (Noztek, West Sussex, UK) was used to extrude the composite PG pellets into filament for 3D printing. The prepared pellets were first fed through the hopper and driven forward by a motor via the rotating screw. Two consecutive heater bands at the end of the extruder, set to 120°C and 100°C, created a temperature gradient to fuse the pellets into a molten state, after which it was extruded through a 3mm nozzle at a slow rate (5 rpm) to allow for the viscous material to flow while maintaining a constant diameter. Extruded filament was then wound into a spool and stored in a cool, dry, sealed container with desiccants to minimize moisture absorption from the environment. Prior to 3D printing, the filaments were dried at 45 °C for 2h to remove any trace moisture.

[0098] A desktop 3D printer, Ultimaker S3 (Dynamism, USA), was used to print custom designed disks, scaffolds, and grafts. An extrusion temperature of 110°C and speed between l-5mm / s was used, with increasing gelatin content corresponding to slower print speeds and higher material flow compensations (120-140%). To circumvent the fixed plate leveling temperature in the Ultimaker firmware, which automatically heats the print core to 200 °C (above the decomposition temperature of gelatin) when leveling the print bed before every run, the print cores were cleaned using a standard hot and cold pull method to remove any gelatin residue in the chamber prior to initializing a print. The filament was then manually loaded after the leveling procedure, when the temperature of the cores dropped to the set temperature (110 °C).

[0099] Mechanical testing was performed using an Instron 5944 uniaxial testing system, following the protocol as previously described

[0026] . Briefly, 5 cm filament segments were cut for tensile testing, and porous scaffolds (10mm x 5 mm) with a rangeof porosities (0-80%) were 3D printed for compression testing. A 2 kN load-cell (Instron Corporation, Norwood, MA) was used, with a preload of IN and a speed of 1% strain / s until 25% strain. The elastic modulus and tensile strength of the materials were extracted from the linear and zero-slope regimes of the stress-strain curves, respectively, and the compressive strength was calculated from the force-extension curves.

[0100] Surface hydrophobicity was characterized using a contact angle goniometer (Rame-Hart 290, Rame-Hart Instrument Company, USA). Solid disks (7mm x 0.7mm) of different ratios of gelatin were 3D printed prior to the experiment. A single droplet (4 L) was deposited at the center of each disk, and image processing (DROPImage, ImageJ) was used to obtain and calculate the surface contact angles. Measurements were repeated ten times for each material group, and left and right contact angles were averaged to obtain the mean contact angle for each material composition.

[0101] A thermogravimetric analyzer (TGA) instrument was used to assess the material composition and thermosensitivity of the PCL-Gel blends (TA 550, TA Instruments, USA). Briefly, disk samples of 5mm diameter and 1mm height were 3D printed for each material composition and placed into ceramic pans (alumina crucible, I OOpL, TA Instruments, USA). The samples were heated to a maximum temperature of 550 °C with a 20°C per minute ramp rate to assess the thermal degradation of each material blend.

[0102] A FTIR spectrometer (Nicolet iS50) was used to assess the spectra of the composite PCL-gelatin materials. Attenuated total reflection (ATR) measurements were conducted at 32 scans per sample with a resolution of 4 cm'1(4000 cm'1- 500 cm'1). Disk samples (5mm diameter, 1mm height) were 3D printed for each composite material and directly placed above the built-in diamond ATR window prior to measurements. Gelatin type A (Sigma Aldrich) in its compressed powder form was included as a positive control. The resulting spectra were assessed for group and fingerprint absorption peaks and plotted using Origin (OriginPro, Origin Labs, USA).Degradation Kinetics

[0103] An accelerated degradation protocol

[0010]

[0026] was used to quantify the differences in degradation rate among PG material blends in an alkaline environment. A long-term degradation study was also conducted, wherein segments were immersed in lx PBS (Gibco, USA) to discern differences in material degradation in a neutral pH environment, as in vivo. Filament segments (5 cm) were weighed for dry mass before immersing in either alkaline solution (5N NaOH) or PBS and incubated at 37°C. Specimens were removed at designated time points, washed 3 times with deionized (DI) water to remove any surface salts, dabbed with tissues (Kimwipes, Kimberly- Clark, USA), and dried in the oven at 45 °C to remove residual moisture content. The remaining dry mass was then measured, and degradation was quantified by calculating the percentage of remaining mass at each time point. For samples in alkaline solution, degradation measurements were discontinued after the samples could no longer be collected due to complete disassembly from hydrolysis.Scanning Electron Microscopy

[0104] A scanning electron microscope (Apreo S LoVac, Thermo Fisher Scientific) was used to characterize the surface and radial morphology of the bulk materials. Filament cross-sections and longitudinal surface segments were sliced with a microtome blade and mounted onto specimen mounts (25.4mm) using conductive silver paint (PELCO, Ted Pella Inc). Disks seeded with cells were first washed three times with PBS and then fixed with 2% glutaraldehyde (Electron Microscopy Sciences) at room temperature for 1 hour. Graded dehydration (Ethanol, 50% - 100%) was conducted prior to immersing the disks in hexamethyldisilazide (HMDS, Electron Microscopy Sciences) to completely dry the samples. Gold and palladium (Au / Pd 60:40) sputter coating (SNSF, Stanford, USA) was used to coat all samples prior to imaging with SEM. Images were collected at an acceleration voltage of 3 kV and current of 50 pA at magnifications between 5,000x to 50,000x.In Vitro Biological Study

[0105] Human mesenchymal stem cells (hMSCs) were cultured in Dulbecco’s Modified Eagle Medium (DMEM, Gibco, USA) media with 1% FBS in a 750 cm2flask and cultured until confluent, then trypsinized and counted. To quickly assess whether gelatin retains its bioactivity after repeated exposure to high temperatures, bulk material was heat cycled in an oven 3 times for 1 hour at 100°C prior to soakingin DMEM for 24 hours. This primed media was then collected, mixed with human mesenchymal stem cells (hMSCs) at a density of 10,000 cells / mL, and seeded onto an ultra low adhesion well plate (Corning Costar, Fisher Scientific, USA). Cell attachment was observed the next day using a brightfield microscope at lOx magnification.

[0106] To then assess material cell toxicity, a primed media study using filament segments (5mm) was conducted. hMSCs were seeded in a 24-well plate (Coming Costar, Fisher Scientific, USA) at a density of 10,000. Transwell inserts with 0.4 pm pore size (Corning Transwells, Fisher Scientific, USA) were inserted with the filament segments, and 1 mL of media was added to fully submerge the filament segments within the Transwells. The hMSCs were cultured for 8 days, and cells were imaged at designated time points to assess morphological changes and proliferation rate.

[0107] To directly assess material-dependent cell proliferation, thin disks (14 mm x 0.75mm) were 3D printed for each material composition (hPCL, PG10, PG20, PG30, and PGT-712, respectively). The printed disks were then treated with 5N NaOH (Ricca Chemicals) for 1 hour at room temperature to hydrophilize the surface before washing three times with PBS (lx, Gibco). All samples were sterilized using 70% ethanol for 30 minutes before washing thrice with sterile PBS and drying overnight in a ventilated biosafety cabinet. For cell seeding, hMSCs were seeded on the surface at a density of 10,000 cells and cultured for 1 week in FBS-containing DMEM media. At each time point, cell proliferation was quantified using an MTS assay (OD = 490 nm) after a 2-hour incubation period. To assess MSC differentiation, disks were prepared as above, and an alkaline phosphatase (ALP) assay (Abeam, USA) was conducted following the manufacturer’s protocol at the designated time points. Samples were washed thrice with PBS and dried prior to imaging. The level of ALP staining was quantified using ImageJ by calculating the weighted average of blue pixels per pixel intensity (on a scale of 0 to 255) and normalizing by the area of the disk.Subcutaneous Implantation of Filament in a Rat Model

[0108] Eighteen male 10-11 week-old Sprague-Dawley rats (Kyoto, Japan) were used. Anesthesia was induced and maintained with isoflurane (5% for induction and 2% for maintenance, with 2 L / min oxygen flow). Analgesia was administered by an injection of 0.05 mg / kg of buprenophrine, and antibiotics were administered by enrofloxacin (5 mg / kg). Five material groups were tested: PCL (control), PG10,PG20, PG30, and PGT-712. Samples were prepared by cutting a 2.85mm diameter filament into 10mm segments to represent the bulk material composition, and all samples were terminally sterilized via electron beam (E-Beam) irradiation (25 kGy dosage) prior to implantation. Each rat received five implants (one per material group) by subcutaneous implantation in the dorsal skin. Five unconnected incisions were made in two rows across the spine to create subcutaneous pouches. Each filament sample was implanted in a pouch and then sutured to seal the subcutaneous tissue and skin. The implant volume and in vivo response were evaluated at 3 time points (4 weeks, 12 weeks, and 24 weeks, N=6 rats per time point) after euthanasia and harvesting of the surrounding tissue. The change in volume was calculated by the difference between preoperative and post-euthanasia volumes, as determined by p-CT imaging (SKYSCAN 1275, Broker, USA) at 30 mA, 150 kV, and 10 pixels.

[0109] Postoperative samples were demineralized with EDTA (Sigma Aldrich) and dehydrated in serial concentrations of ethanol. All samples were embedded in paraffin blocks and sliced into 5 pm specimens using a microtome blade (Leica RM26555). Hematoxylin and eosin (H&E) staining and immunostaining (Anti-CD68) was performed on surrounding tissue following the manufacturer’s protocol.Femoral Defect Repair in a Rat Model

[0110] 12 weeks-old male Sprague Dawley rats were purchased from Charles River Laboratories. The rats were anesthetized with 3-5% isoflurane (Fluriso, VetOne, Boise, ID) for induction and 2% for maintenance. Buprenorphine-SR (Img / kg, ZooPharm, Laramie, WY) and cefazolin (25 mg / kg; TCI, Tokyo, Japan) were administrated subcutaneously, and Bupivacaine (2mg / kg) was also injected to the surgical site before surgical procedure. The skin of the surgical site was sterilized with 70% EtOH and povidone-iodine, respectively, three times. A skin incision was made on the lateral side of the left thigh, and skeletal muscle was separated from the femoral diaphyseal periosteum. 8 mm bone defects were created in the middle of femurs, and the residual bone edges were fixed with the external fixator. Then, rats were randomly divided into three material groups: PCL (N=4), PCL-TCP (N=3), or PG30 (N=4). The inserts with a length of 18 mm and diameter of 1 mm were 3D printed (Ultimaker S3, Ultimaker USA) and surface treated with NaOH (5N, Ricca Chemicals, USA) for 1 hour at room temperature, washed thrice with PBS, and UV sterilized prior to implantation into the bone marrow cavity to use as intramedullaryimplants. Carprofen (5 mg / kg; Zoetis, Parsippany, NJ) was administrated subcutaneously to the animals right after surgery and once a day for 3 days. All animal experiments were approved by the Institutional Animal Care and Use Committee (IACUC) of Stanford University following the ARRIVE guideline. Animals were singly housed after operation. At post-operative day (POD) 28, the external fixators and pins were removed from femurs, and the femurs were harvested after euthanasia using CO2. The femoral specimens were fixed in 10% neutral buffered formalin at 4 °C for 48 hours, then transferred to 70% ethanol for further analysis. To evaluate bone formation at defect sites, X-rays (Faxitron Bioptics, Tucson, Arizona, USA) were taken at 22kV on POD 1, 7, 14, 21, and 28. p-CT images were taken on POD28 for the quantification of bone formation at the bone defect sites. Femoral specimens were tested with a three-point bending test (Instron 5944, Norwood, MA) to evaluate the mechanical properties. Histological samples were embedded in paraffin blocks and sliced into 5 pm specimens using a microtome blade (Leica RM26555) and stained with hematoxylin and eosin (H&E) and Masson’s Trichrome to evaluate cell morphology and new bone formation.Femoral Defect Repair in a Canine Model

[0111] To fabricate the graft, computer-aided design (Onshape) was first used to create a cylindrical model of 1cm in diameter and 3cm in length with a center channel of 4mm, which was then exported into a Standard Tessellation Language (STL) file format. The porosity was set to 80% (Ultimaker Cura) to allow space for tissue infiltration, and the grafts were printed via molten material extrusion (Ultimaker S3, USA). The grafts were treated with alkaline solution (5N NaOH) for 1 hour to further hydrophilize and remove any trace debris on the surface, washed thrice with lx PBS, and dried overnight. All grafts were terminally sterilized with E-Beam irradiation prior to implantation, and a total of three groups were tested: disease control (no implant), PG30, and PG30 + bone morphogenetic protein-2 (BMP-2). BMP-2 was added by first immersing the implant into a solution of fibrinogen prior to manually injecting BMP-2 (150 pg) onto the surface. Micro-computed tomography (p-CT) was used to assess the vascular volume and implant volume before and after implantation.Results3D Printing

[0112] Since PCL is the only extrudable material component, increasing gelatin and tricalcium phosphate composition resulted in lower extrudability, which had to be compensated for in the print parameters, primarily by lowering the print speed and increasing the material flow rate of the filament during the print. Table 1 summarizes the key printing parameters across PCL-based filaments, irrespective of the printed construct dimensions. PCL-TCP at an 80 / 20 ratio is included as a base-line, as this has been characterized as a bone tissue engineering material candidate in prior work

[0026] . 3D printed PCL-Gelatin scaffolds with varying porosities are provided in FIG 2.Table 1: PCL-gelatin printing parameters for molten material extrusion (MME).Instron Testing

[0113] Uniaxial tensile testing of filament segments (5 cm) revealed that increasing gelatin as well as P-TCP content increased the stiffness (Young’s modulus) of the composite material, which agrees with the literature

[0026]

[0018] . Increasing P-TCP also increases the brittleness of the composite material, thus lowering the maximum tensile force sustained by the material before failure, while increasing the gelatin content increases the toughness of the material, allowing it to sustain a significantly higher peak force (FIG 3). As expected, increasing the porosity of a 3D printed scaffold (disk) from 0% to 80% resulted in a general trend of decreased compressive modulus, as the strut distance, and thus the force sustained by each strut, increases with higher porosities. Interestingly, gelatin-containing groups appear to be more resistant to compressive weakening with increasing porosity compared to PCL alone, which showed a significant drop in the compressive modulus at 80% porosity compared to the PG materials.Scanning Electron Microscopy

[0114] All filaments containing gelatin appeared to have a bumpy surface texture under SEM, and radially sliced cross-sections for all composite groups revealed nano-sized particles embedded within a matrix (FIG 4). On the contrary, the hPCL control group, which contains no gelatin or beta-tricalcium phosphate content, showed some irregular surface asperities as a result of filament extrusion and microtome slicing, but was otherwise smooth and free of any embedded particles. The PCL-TCP group, which does not contain gelatin, showed very irregularly shaped nanoparticles under SEM, whereas the PGT-712 group contained both spherical and irregularly shaped nanoparticles. Since the fabrication and processing steps among groups are the same, this indicates that the spherical nanoparticles observed embedded within the PGContact Angle Goniometer

[0115] Contact angles decreased with increasing gelatin content, with an average contact angle of 82.2°, 81.2°, 66.5°, 51.8°, and 76.8° for hPCL, PG10, PG20, PG30, and PGT-712, respectively. This trend indicates that increasing gelatin improved the hydrophilicity of the composite material compared to PCL alone. On the other hand, the presence of tricalcium phosphate only slightly increased the hydrophilicity of the composite when compared to gelatin alone. FIG 5 shows contact angle measurements.Thermo gravimetric Analysis

[0116] The thermal degradation temperatures of gelatin (212 °C) and PCL (300 °C) are reflected in the material trends (FIG 6). For the materials comprising PCL and gelatin, a steady loss in mass is first observed, corresponding to the thermal degradation of gelatin, before a sharp drop, indicative of the dominant effects of PCL. Pure PCL displayed a single stage thermal degradation at 300 °C, which is lower than prior literature reports (in the range of 350 to 370 °C) but internally consistent with our synthesized PCL and raw material pellets. The single stage thermal degradation of PCL has been reported to be due to hydrophilic ester group breakage and long-chain hydrocarbyl rupture

[0027] , which indicates that the dual stage thermal degradation observed in different PG blends are the combinatory result of PCL and gelatin thermal properties. In addition, for each composite material, the point at which a sharp, PCL- dominant thermal degradation occurs aligns closely with the designed material ratios, with PG10 exhibiting a sharp drop in mass at 90% of its initial mass, PG20 at 80% ofits initial mass, and so on. Furthermore, for the PGT-712 group, which contains 20% TCP, which does not thermally degrade within the temperature range tested, the remaining mass asymptotically dropped to 20% and resembled pure-white TCP nanocrystals.Fourier Transform Infrared Spectroscopy

[0117] The ATR spectra of PG composite materials revealed the expected absorption spectra of PCL and gelatin. At 3311 cm1, corresponding to aliphatic primary amine (N-H) stretching, increasing peak intensity can be observed for increasing gelatin ratios, with PG30 being the highest (FIG 7). Similarly, at 1657 cm1and 1544 cm1, corresponding to amine (N-H) bending and nitro compound (N-O) stretching, PG30 exhibited the highest absorption intensity. When compared to the spectral absorbance of pure gelatin, the peaks of PG composite materials corresponding to amine and amide groups appear to be shifted leftward and narrower, indicating a synergistic effect of gelatin within the material beyond purely physical mixing. The leftward shift in peaks can be interpreted as shortening of N-H bond lengths in the composite materials, whereas the changes in peak shape are indicative of new hydrogen bonding

[0028] .

[0118] The sharp peaks observed at 1721 cm'1for all groups correspond to strong aliphatic ketone (C=O) bond stretching, which is present in the chemical structure of both PCL and gelatin. The peak intensity is the highest in the control group (PCL), which has a carbon backbone and repeating monomer units containing the C=O bond. This change in relative peak intensities appears to indicate that the presence of gelatin in the composite materials is associated with new spectral features that change the absorption spectrum of the material as a whole

[0029]

[0030] .

[0119] The accelerated degradation kinetics of the composite materials in alkaline conditions (NaOH) show a clear difference among the groups, with PG30 exhibiting the fastest degradation rate at 14h, followed by PG20, PGT-712, and PG10 at 24h, 30h, and 41h. These time points reflect the stage at which the filaments radially disassembled to a point at which reliable mass measurements could no longer be made. The control group, PCL, showed no significant degradation even in alkaline conditions within the time frame of 2 days and extending beyond to 1 week (data not shown). The rapid degradation appears to be directly proportional to the gelatincontent, indicating that increasing the gelatin content of the composite significantly increases the degradation rate of the hybrid material. On the other hand, the addition of 0-TCP in the PGT-712 group appears to only slightly increase the accelerated degradation rate when compared to PG 10, which contains the same percentage of gelatin. When compared to PG30, which contains the same mass ratio of PCL, PGT-712 exhibits a significantly slower degradation kinetics profile.

[0120] The difference in degradation rates among material groups was much less pronounced in the long-term in vitro PBS study at physiological pH (7.4). There was no significant difference among the control group, PCL, PG10, and PG20 groups, which only exhibited 1-2% change in mass over 6 weeks. PG30 experienced a mass loss of 7% at the end of 42 days, which was significantly different from the other groups, indicating that even in the absence of enzymatic activity and dynamic flow conditions in vivo, increasing the gelatin content to 30% mass ratio significantly increases the degradation rate over the span of 6 weeks (FIG 8).Bioactivity of Gelatin After Heat Cycling

[0121] The hMSCs primed in media exposed to hPCL and then seeded onto an ultralow adhesion well plate formed spherical aggregates (FIG 9) and remained floating in the cell culture media after 1 day. By contrast, cells that were primed with PG30 media attached to the low-adhesion well plate and exhibited clear changes in cell morphology, showing cell attachment, elongation, and network formation across the entire well plate. The cells primed with media containing dissolved gelatin powder, on the other hand, were also able to attach to the low-adhesion well plate, but with the presence of some cell aggregates and less cell elongation. These results indicate the retention of gelatin bioactivity even after repeated exposure to high temperatures (100°C), with a notable difference in the efficacy of PG-eluted gelatin nanoparticles in promoting early cell attachment and growth. hMSC Proliferation in PCL- Gelatin Primed Media

[0122] Despite equal cell seeding on Day 0, the density of hMSCs appears to be slightly higher in the PG30 group after 1 day of culture and becomes pronouncedly higher at subsequent timepoints. This indicates that cells indirectly exposed to PG30 proliferate at a faster rate compared to those primed with the PCL control. This trend continues over the span of 1 week, with the PG30 group reaching full confluence by Day 4, while the PCL group does not become confluent until Day 6. After reachingconfluence, the contact-inhibited cells begin to detach at the edges of the wells, as shown on Day 6 for the PG30 group and Day 8 for the PCL group (FIG 10). hMSC Proliferation on 3D Printed PCL-Gelatin Disks

[0123] All groups showed an increase in cell metabolic activity over the span of 1 week, with a comparable absorbance reading on Day 0 of cell seeding. All groups except PG10 and PG30 show a significant drop in cell metabolic activity after 1 day of culture, indicating that some seeded cells may have detached in the other groups during periodic media changes. However, all groups proliferate and show in increase in metabolic activity by Day 5, with the PG30 group showing the highest increase in cell proliferation through the 1 week period. Additionally, increasing gelatin content in PG materials appears to be positively correlated with increased proliferation, whereas the addition of only TCP slightly increases cell proliferation when compared to the control (PCL). The cell morphological differences are clear through the SEM images taken of samples after 7 days of culture. While hMSCs on PCL are attached, they appear to be flat, with a noticeable lack of cell protrusions and migration from the initial location of cell seeding compared to the other groups. On the other hand, the PG20, PG30, PGT-712, and PCL-TCP groups all exhibit a high number of filopodic extensions and surface particles resembling extracellular vesicles (EVs), indicating an intracellular response to the nanosized particles in the materials (gelatin and P-TCP). This finding agrees with other reports indicating that nanoparticles can modulate extracellular vesicle production in mesenchymal stem cells

[0031] . hMSC Differentiation on 3D Printed PCL-Gelatin Disks

[0124] The PCL and PCL-TCP group exhibited no ALP activity on Day 1 of culture, which is as expected. Interestingly, however, the PG30 and PGT-712 groups, which showed the highest density of extra-cellular vesicle-like structures under SEM, had low levels of ALP staining even on Day 1 of culture in the absence of osteogenic medium. On Day 14, the PGT-712 group had a significant increase in ALP activity compared to the other groups, indicating that a synergistic effect of gelatin and [3- TCP nanoparticles can stimulate higher hMSC differentiation compared to PCL, PCL- Gelatin, or PCL-TCP alone, even in the absence of osteogenic medium or other biological cues. FIG 11 shows an experimental setup. FIG 12 shows alkaline phosphatase staining on 3D printed disks.

[0125] No rats developed postoperative complications in the in vivo rat experiments, although one of the six rats assigned to 24-week followup died before reaching 24 weeks. From the p-CT calculations comparing preoperative and postoperative implants, PCL decreased in volume at 4 weeks (-0.77%), whereas all four other groups showed an increase in volume, with PG10: +3.53%, PG20: +5.5%, PG30: +6.61%, PGT-712: +7.01%) (FIG 13). At 4 weeks, there was a significant difference between the rate of change in volume of PCL and the other four samples, and the volume change rate of PG10 was significantly different from that of PG30 and PGT-712.

[0126] At 12 weeks, the control volume was further reduced, and all four new samples showed a volume increase compared to preoperative levels. However, the rate of increase in the four new samples was lower than that at 4 weeks, except for PG20. At 12 weeks, there was a significant difference between the rate of change in volume of the PCL control and that of the other four samples. The rate of change in volume of PG10 was significantly different from that of PG20 and PGT-712, and the rate of change in volume of PG30 was significantly different from that of PG20 and PGT-712. At 24 weeks, control PCL decreased further, and the four other groups all showed an increase compared to preoperative levels. However, the rate of increase in PG10, PG20, PG30, and PGT-712 was lower than that at 12 weeks.

[0127] At 24 weeks, there was a significant difference between the volume change of the control and other four samples. The volume change of PG 10 was significantly different from that of PG20 and PGT-712, and the volume change of PG30 was significantly different from that of PG20 and PGT-712. In summary, the volume of PCL gradually decreased, whereas the volume of the PG10, PG20, PG30, and PGT- 712 increased by week 4 before gradually decreasing, with the exception of PG20, which increased slightly from +5.5% at week 4 to +5.63% at week 12. This can be attributed to the gelatin component of the PG materials, which initially absorbs water and expands before degrading in vivo, resulting in a decrease in volume. Among the four PG groups, PG30 showed a particularly large decrease in volume after the initial expansion, which is promising for its use as a bone tissue biomaterial.

[0128] FIG 14 shows histological sections of H&E stained tissue surrounding the subcutaneous filament implants after 24 weeks. All of the histological sections show anencapsulating layer of tissue surrounding the subcutaneous implant. Although this may seem to be a negative response, indicating fibrosis, studies have shown that fibrous encapsulation of implants is the first step toward healing and osseointegration

[0032] , Furthermore, immunostaining of the tissues indicated low levels of inflammation and macrophage count. The increase in the density and thickness of tissue encapsulation corresponding to increasing concentrations of gelatin in the PG materials could therefore be interpreted as an improvement in cell attachment and the recruitment of localized cells to the implant, which can be beneficial for bone healing applications, wherein osteoconduction and native tissue infiltration is a key consideration in biomaterial selection.Femoral Defect Repair in a Rat Model

[0129] The X-ray images of the live rats in FIG 15, which visualize the hard tissue formation at the femoral defect region, did not show a significant difference in bone volume between the three groups on Day 28, although it appears that the PCL-TCP and PG30 groups formed a higher volume of new bone tissue at an earlier time point than the PCL group (Day 14). The p-CT images taken of the femur explants after animal sacrifice (FIG 16(a)) also showed a comparable level of tissue formation within the defect region, and the quantified bone volume versus tissue volume ratio yielded no significant difference between the groups for length of time of this study. However, as shown in FIG 16(b), there appeared to be a significant difference between the mechanical properties of the bone from the PG30 group compared to the PCL-TCP group, as evaluated by maximum load, Young’s Modulus, and Energy Absorption, as well as a positive trend when comparing PG30 with the PCL control. This seems to indicate that while the bone volume fraction is comparable across material groups at the 4- week mark, the quality of bone, particularly hard bone tissue imparting native bone tissue with its mechanical strength, may be higher in the PG30 group.

[0130] The hematoxylin and eosin and Masson’s trichrome staining showed no observable signs of inflammatory response for any of the implanted materials. The PG30 implant appeared to have a higher density of newly formed bone (NB) compared to the two controls, PCL and PCL-TCP, as shown by the dark blue color in FIG 17, which is distinct from the dense, mature intact bone surrounding the defect site. The yellow arrows indicate regions at the periphery of mature and newly formed bone that contain clusters of cells with a faint blue staining, which can he interpretedas preosteoblast or osteoid cells. These results are promising, as they support the in vitro observation that gelatin eluted from PG materials promotes cell proliferation and differentiation.Femoral Defect Repair in a Canine Model

[0131] A critical size femoral bone defect canine model (3cm) was used to evaluate the in vivo response of PCL-Gelatin in a hard tissue environment, and PG30 was selected as the model material candidate for assessing bone healing, as it exhibits fast degradation and high biocompatibility in vitro and in small animal model evaluations. When comparing the three groups, surgical control (no implant), PG30 only, and PG30 with BMP-2, we can see a clear increase in vascular volume and a positive trend in bone volume in the PG30 groups. Furthermore, we can see signs of angiogenesis and vascular network formation at the center of the implant for the PG30 + BMP -2 group. This is promising, as one of the key limitations of large bone defect repair is preventing necrosis at the center of the implant, which can occur due to loss of blood supply, oxygen, and nutrients from lack of vascularity. As angiogenesis is the first step toward bone healing, these results are a promising indication of PG30 as a bone tissue engineering material candidate for large bone defects.Method

[0132] A 30mm bone defect was created in the femur of six female beagle dogs, and the segment was fixed with a peek plate. Four legs were inserted without any scaffold in the bone defect (Empty group), four legs were inserted with PG scaffold (PG group, 70 wt%PCL-30 wt%gelatin), and four legs were treated with PG and rhBMP-2 ( 150pg) (PG+BMP-2 group) for three-group comparison. Eight weeks after surgery, the amount of neovascularization and osteogenesis were examined using micro-CT and histological evaluation.Results

[0133] The PG scaffolds were 3.6 + 0.1 mm in lumen, 10.8 + 0.3 mm in outer diameter, and 30.4 + 0.3 mm in length, with a porosity of 73.7 + 6.1%. The amount of new bone formation in the defect area was 145.4+109.5 mm3in the Empty group, 177.7+57.9 mm3in the PG group, and 317.1+112.9 mm3in the PG+BMP-2 group. The volume and ratio of vessels inside the scaffold was 26.9+13.9mm3(l.5+0.7%) in the Empty group, 51.5+20.5mm3(2.3+0.9%) in the PG group, and 50.9+18.3mm3(2.1+0.8%) in the PG+BMP-2 group. The volume and ratio of vesselsoutside the scaffold within a diameter of 0 15mm was 33.3±16.3mm3(l.5+0.5%) in the Empty group, 84.2±20.9mm3(3.5+0.7%) in the PG group, and87.0+58.0mm3(3.5+0.2%) in the PG+BMP-2 group. The micro-CT evaluation showed that PG group and PG+BMP-2 groups tended to promote angiogenesis and osteogenesis compared to the Empty group. The reduction rate of the scaffold width was 4.4+2.6 % in the PG group and 4.2+1.2 % in the PG+BMP-2 group.Conclusion

[0134] Although further addition of growth factor should be considered to obtain complete bone healing, its ability to promote angiogenesis and osteogenesis was confirmed in this study. PG scaffold has the potential to be used for the large defect in the future with further modification.Introduction

[0135] Repairing extensive bone defects resulting from significant trauma, inadequate fracture healing, or surgical tumor excision remains a significant clinical challenge. Every year, millions of bone graft procedures are performed worldwide, many of which involve extensive bone grafts to address these complex conditions [l a].

[0136] The treatment of large bone defects requires not only substantial structural support but also the regeneration of bone tissue that can integrate effectively with the patient's existing bone. Conventional methods, such as autologous and allogeneic bone grafts, are widely used but come with significant drawbacks. The use of autologous bone grafts is limited by donor site morbidity, inadequate supply, and variable quality of the graft material [2a]. Similarly, allogeneic bone grafts pose risks such as immune reactions and disease transmission. These limitations underscore the necessity for developing alternative strategies to effectively repair large bone defects.

[0137] Bone tissue engineering (TE) has surfaced as a promising alternative for overcoming the limitations of current treatments. TE holds immense promise in revolutionizing modern medicine by addressing critical healthcare challenges such as organ shortages for transplantation and the limitations of conventional medical treatments, through the creation of biological substitutes that restore, maintain, or improve tissue function. Polycaprolactone (PCL), an FDA-approved biodegradable polyester, is a key material in TE, renowned for its favorable mechanical properties, biocompatibility, and processability [3a], Widely utilized in manufacturing 3Dscaffolds for bone TE, PCL is valued for its cost-effectiveness, slow degradation rate, high load-bearing capacity, and versatility in shaping.

[0138] Among the techniques in TE, 3D bioprinting stands out for its ability to precisely control scaffold architecture and internal porous structure, which are crucial for cell adhesion, proliferation, and tissue ingrowth in bone regeneration applications. This technology allows for customizable scaffold designs with resolutions ranging from micrometers to millimeters, offering tailored solutions to match the complex geometries of bone defects and optimize therapeutic outcomes in clinical settings [4a].

[0139] Recombinant human bone morphogenetic protein 2 (rhBMP-2), known for its osteoanabolic properties, has been clinically employed via collagen sponge delivery to stimulate osteoblast differentiation in treatments such as lumbar fusion and tibial diaphyseal fracture repair [5a] [6a]. rhBMP-2 is a valuable tool in orthopedic and reconstructive surgeries to enhance bone healing and regeneration.

[0140] Researchers have endeavored to harness the potential for self-repair in their designs of systems aimed at healing larger defects. Previous research has highlighted the efficacy of rhBMP-2-loaded PCL scaffolds in promoting osteogenesis and angiogenesis in rat femoral defects, demonstrating significant promotion of both processes [7a].

[0141] Utilizing a large animal model (dog femoral defects), this study assesses the effectiveness of rhBMP-2-loaded 3D printed PG scaffolds in promoting vascularity and facilitating bone defect repair. The trabecular structure of canine bone closely resembles that of human trabecular bone, making it an ideal model for translational research in orthopedic applications [8a].Methods1 Scaffold preparation1.1 Material synthesis

[0142] Polycaprolactone (PCL, 80,000 Mn, Sigma Aldrich, USA) and Gelatin type A derived from porcine skin (Sigma Aldrich, USA) were weighed at a 70:30 weight ratio. The materials were dissolved in 3-hexafluoro-2-propanol (HF1P, 99% purity, Synquest Laboratories, USA) to a final concentration of 5% (w / v). The solution was incubated in a shaker at 37°C and 80 rpm for 24 hours to ensure complete dissolution. To minimize phase separation between PCL and Gelatin, glacial acetic acid (Sigma Aldrich) was added the next day to the solution at a concentration of 0.05% (v / v).This adjustment in pH was crucial to improve the miscibility of the hydrophobic PCL and hydrophilic Gelatin components. The homogeneous solution was poured into low-form polytetrafluoroethylene (PTFE) evaporation dishes (Fisher Scientific, USA) and maintained at 70°C under continuous mixing to facilitate even casting. The solvent was allowed to evaporate under a chemical fume hood overnight.Subsequently, the dishes were removed from heat and left to evaporate for an additional 72 hours at room temperature.

[0143] The dried composite sheets were manually removed from the PTFE dishes and cut into pellets approximately 3 mm in diameter using stainless steel blades. These pellets were further air-dried in the fume hood for 24 hours to ensure the removal of residual solvents. To eliminate any residual acetate salts, the pellets were repeatedly washed in cold water (4°C) until neutral pH was achieved, followed by deep freezing at -80°C and lyophilization for 48 hours.1.2 Filament Fabrication

[0144] The dried and purified PCL / Gelatin pellets were fed into a filament extruder equipped with a stainless-steel barrel and a custom-sized screw designed to accommodate particles smaller than 5 mm (Noztek, West Sussex, UK). The extruder was set to create a temperature gradient with heater bands at 120°C and 100°C, fusing the pellets into a molten state. The composite was extruded through a 3 mm nozzle at a slow rate of 5 rpm to maintain consistent filament diameter. The resulting filament was wound into spools and stored in a cool, dry container with desiccants to prevent moisture absorption.1.3 3D Printing

[0145] Prior to 3D printing, the filament spools were dried at 45°C for 2 hours to remove any trace moisture. A desktop 3D printer (Ultimaker S3, Dynamism, USA) was utilized for printing custom-designed scaffolds and constructs. The extrusion temperature was set at 110°C, and the print speed was adjusted between 1-5 mm / s, with slower speeds corresponding to higher Gelatin content to ensure adequate material flow and print fidelity.2 Animal surgical procedure2.1 Preparation of Scaffolds

[0146] To incorporate rhBMP-2 into the scaffolds, the following procedure was used: First, the scaffolds were placed in a 1 mg / ml fibrinogen solution and incubated at37°C for 24 hours, followed by three washes with sterile PBS. Subsequently, they were immersed in a 4U / mL thrombin solution under the same conditions and washed again with PBS three times. Lastly, each scaffold was treated with 0.3 ml of a 0.5 mg / ml rhBMP-2 solution, resulting in a total of 150 pg of rhBMP-2. The treated scaffolds were then stored overnight at 37 °C before being utilized in the animal experiments the following day.2.2 Animal experiments in vivo

[0147] Six female beagles, aged 10-13 months and weighing 10-11 kg, were used in this study to analyze bone defect healing and vascular formation across three experimental groups, involving a total of 12 legs. Group 1 had no scaffold in the bone defect (Empty group, n=4); Group 2 had PG scaffold alone (PG group, n=4); and Group 3 had rhBMP-2 (150 pg) loaded PG scaffold (PG+BMP-2 group, n=4). Each dog had both hind legs operated and each hind leg was randomly assigned to one of three groups.

[0148] This experiment was approved by the Institutional Animal Care and Use Committee of Kyoto University (Approval number: Med Kyo 21557).2.3 Surgical Procedure

[0149] Anesthesia was induced with sevoflurane (2%) following subcutaneous injections of seractol (2%, 1 ml), dolmicum (5 mg / 1 ml), and atropine (0.25 mg / 0.5 ml). After intubation, anesthesia was maintained with sevoflurane (2%).Postoperatively, 5 mg / kg of enro floxacin was administered for antibiotic therapy.

[0150] A 10 cm skin incision was made via a lateral femoral approach. The fascia lata over the vastus lateralis was cut, and the vastus lateralis muscle was retracted to reveal the femur. A 90 mm polyether ether ketone (PEEK) plate was temporarily fixed to the anterior femur with 2.4 mm Kirschner wires and then secured with six 2.7 mm screws. A 30 mm bone defect was created in the center of the plate using a rotating diamond-coated drill bar. In the PCL and PCL+BMP-2 groups, the scaffold was inserted into the bone defect, which was created slightly smaller than the scaffold to allow for a press-fit stabilization within the defect (FIG 19(a)). Finally, the fascia and skin were sutured.3 Evaluation3.1 Radiographic Evaluation

[0151] X-rays of the femur were captured at 0, 4, and 8 weeks after surgery using an X-ray device (BRANSIST- Alexa, Shimadzu, Japan) to evaluate bone formation and detect adverse events such as scaffold displacement.

[0152] The structure of the scaffolds was examined before implantation using a micro-computed tomography (pCT) scan (Skyscan 1275, Bruker, Billerica, MA, USA). This scan was performed without a filter, with a pixel size of 25 pm, a voltage of 65 kVp, and a current of 75 p A. Hindlimb angiogenesis was assessed at 8 weeks after surgery via microangiography. An 18-gauge catheter was inserted into the abdominal aorta, and the aorta was flushed with 1000 ml of heparin solution (30 U / ml). Under physiological pressure, a polymerizing contrast agent (Microfil MV- 122 Yellow, Flow Tech, Carver, MA) was administered through the abdominal aorta [7a]. Following this, the thigh was extracted and subjected to a pCT scan to measure bone and blood vessel volume. This scan was performed without a filter, with a pixel size of 50 pm, a voltage of 65 kVp, and a current of 75 pA.

[0153] A cylindrical region of interest (ROI) with a diameter of 010mm was set concentrically over the defect site, and the vascular volume (VV) of the entire defect area were measured.

[0154] To differentiate regional differences in vessel volume (VV), two ROIs were defined. Two cylindrical ROI (one with a diameter of 010mm and the other with a diameter of 015 mm) were placed over the defect site along the scaffolds. The ectopic VV were quantified by the 010 mm ROI from the 015 mm ROI. The vessel volume percentages for each region were calculated by dividing the corresponding volume by the total volume (TV). And a cylindrical ROI of 010 mm diameter in a 3 cm defect was divided into distal, middle, and proximal sections at every 1 cm height to measure the percentage of vessels occupied. The volume of new bone formation (BV) was quantified as the total amount of newly formed bone between the original defect sites. The precise location of the original defect was identified based on the femur's shape and the position of the scaffold.3.3 Histology

[0155] Samples containing scaffolds were fixed in 10% phosphate-buffered formalin for 7 days, dehydrated through a series of ethanol concentrations, and embedded in polyester resin. Sections (500 mm) were cut parallel to the scaffold’s axis with a band saw and ground to a thickness of 50-70 mm using a grinding slide device (BS-300CP- A: EXAKT, MEIWAFOSIS, Tokyo, Japan). The samples were stained with Van Gieson’s picrofuchsin and Stevenel's blue and and examined using a transmitted light microscope (BZ-X710, KEYENCE, Osaka Japan)3.4 The scaffold reduction rate

[0156] The rate of scaffold degradation was determined as the reduction rate of the scaffold width at 8 weeks using the following formula:Postoperative scaffold width was measured using postoperative resin- fixed sections without staining. The length of the scaffold running laterally on the resin section was measured (FIG 19(b)). Preoperative scaffold width was measured using pCT images, with the corresponding length on the preoperative pCT. The average of five locations was used for the scaffold reduction rate.3.5 Statistical Analysis

[0157] All data were analyzed using JMP Pro 15.2 software. Mean differences were determined using the Kruskal-Wallis test, followed by the post hoc Steel-Dwass test for pairwise comparisons. A p-value of <0.05 was considered statistically significant. Data are reported as mean ± standard deviation (SD).Results1 The structure of the scaffold

[0158] Based on image processing, the completed 3D printed scaffolds measured 30.4±0.3mm in height, 10.8±0.3mm in outer diameter, 3.6+0.1mm in the inner diameter, and the struct size was 393.3+11.4pm, and the porosity of the scaffold was 73.7+6.1 % (FIG 18(b)). SEM examination revealed that PG scaffold had smooth surface and macropores (FIG 18(c)).2 New bone and vessel

[0159] pCT images were showed at FIG 20. In the PG and PG +BMP-2 group, vessels entering the scaffold structure with invasion from surrounding tissueobserved. Vessels were also observed entering the scaffold lumen from cancellous bone. The amount of new bone formation in the defect area was 145.4+109.5 mm3in the Empty group, 177.7+57.9 mm3in the PG group, and 317.1+112.9 mm3in the PG+BMP-2 group (FIG 22(a)). While PG+BMP-2 group had a larger amount of bone formation, there were no significant differences observed between any two groups (p=0.25 for Empty group vs. PG group, p=0.25 for Empty group vs. PG+BMP-2 group, p=0.75 for PG group vs. PG+BMP-2 group). New bone formation within the defect was observed entering the scaffold tunnels and surrounding the scaffolds in the PG and PG+BMP-2 groups. In contrast, the Empty group exhibited disorganized new bone and vessel formation without clear directionality.

[0160] Histological images presented in FIG 21 showed that in the Empty group, the defect was covered with soft tissue and exhibited disorganized new bone formation. In contrast, the PG and PG+BMP-2 groups demonstrated new bone growth entering the scaffold lumen and surrounding the scaffold structure. Within the scaffold, vessels filled with microfil were observed throughout in the PG and PG+BMP-2 groups.

[0161] FIG 22 shows the results of quantitative evaluation of new bone and new blood vessels at 8 weeks after surgery. The volume and ratio of vessels inside the scaffold was 26.9+13.9mm3(l.5+0.7%) in the Empty group,51.5+20.5mm3(2.3+0.9%) in the PG group, and 50.9+18.3mm3(2.1+0.8%) in the PG+BMP-2 group (FIG 22(b)). There were no significant differences observed between any two groups. Excluding scaffold volume, the ratio of vessels inside the scaffold was 3.5+0.7% in the PG group and 3.5+0.2% in the PG+BMP-2 group (FIG 22(c)). No significant differences were observed between the PCL and PG+BMP-2 groups (p=0.15 for Empty group vs. PG group, p=0.15 for Empty group vs. PG+BMP-2 group, p=1.0 for PG group vs. PG+BMP-2 group). The volume and ratio of vessels outside the scaffold was 33.3+16.3mm3(l.5+0.5%) in the Empty group, 84.2+20.9mm3(3.5+0.7%) in the PG group, and 87.0+58.0mm3(3.5+0.2%) in the PG+BMP-2 group (FIG 22(d)). Although PCL group tended to have a larger amount of vessel than Empty group, no significant differences were observed between any two groups (p=0.08 for Empty group vs. PG group, p=0.25 for Empty group vs. PG+BMP-2 group, p=0.90 for PG group vs. PG+BMP-2 group). The reduction rate of scaffold width was 4.4+2.6% in the PG group and 4.2+1.2% in the PG+BMP-2 group, with no significant difference (FIG 22(e)). The mean percentage of vessels at eachsite is shown in FIG 21(f). In the Empty and PG groups, there was roughly equal angiogenesis at each site; in the PG+BMP2 group, there was a trend toward more vessels on the proximal side, but the difference was not significant.Discussion

[0162] In the present study, relatively larger amount of angiogenesis and bone formation was observed in PG group than in Empty group in critical sized segmental bone defect model in dog femur. The addition of small amount of BMP-2 used in this study did not enhance vessel formation and bone healing.

[0163] The increasing interest in PCL for scaffold fabrication has driven the development of intricate and spatially accurate 3D-printed PCL scaffolds for TE applications. PCL's biocompatibility, biodegradability, and excellent mechanical properties make it an ideal candidate for bone, cartilage, skin, and cardiovascular TE [9a]. Its cost-effectiveness, thermal stability, and versatility in forming different structures further enhance its suitability for creating scaffolds that support tissue growth and regeneration [10a],

[0164] The design and construction of scaffolds play a crucial role in bone TE. 3D printing technology allows precise control over scaffold architecture and internal porous structure, facilitating cell adhesion, proliferation, and tissue ingrowth.Refining the scaffold's microstructure and biochemical environment could overcome current limitations and achieve more effective bone regeneration. Investigating the long-term degradation behavior of PCL and its interaction with biological tissues will provide valuable insights for developing next-generation scaffolds tailored for clinical applications [I la].

[0165] PCL has also been reported to be effective as a cell scaffold. Shor et al. reported that fetal bovine osteoblasts cultured on a 3D porous PCL scaffold exhibited favorable outcomes in cell viability, proliferation, differentiation and calcium content, which suggested the suitability and differentiation inducing ability of the PCL [12a]. Additionally, Shin et al. found that nanofibrous PCL scaffolds aid in angiogenesis in vivo and promote bone formation [13a]. But PCL alone presents certain limitations in vivo due to its hydrophobicity, which adversely affects cell adhesion and proliferation. PCL can be easily blended with various materials to enhance its osteoinductive and osteoconductive properties, proving successful in overcomingthese limitations and enhancing its performance in biological application [14a] [15a] [16a],

[0166] Several attempts have been made to enhance cell growth and differentiation in PCL scaffolds by incorporating signaling factors such as BMP-2. Numerous studies have demonstrated that BMP-2 can effectively promote both bone formation and angiogenesis. Kawai et al. showed that the addition of rhBMP-2 and vascular bundles to PCL scaffolds significantly increased vascularization and bone formation [7a]. BMP-2 stimulates angiogenesis in bone defects by promoting human umbilical vein endothelial cell proliferation and migration [17a] [18a]. Moreover, immobilized BMP-2 on scaffolds has been proven to significantly enhance osteoblast proliferation and osteogenic differentiation [19a]. Vaquette et al. reported that BMP-2-loaded PCL gels augmented bone formation and bone-binding capacity [20a]. Neovascularization plays a crucial role in bone regeneration, with new vessels closely integrated with the forming bone. The scaffold acts as a framework for both blood vessels and cells, facilitating the creation of internal microstructures.

[0167] In this study, 150 pg of rhBMP-2 was used. This is a relatively small amount that is generally used for bone defect models [21a] [22a] [23 a]. BMP-2 stimulates osteogenesis in a dose-dependent manner, and 150 pg was chosen based on reports indicating that this was the lowest effective dose [23a]. However, it is possible that the PG group and the PG+BMP-2 group did not differ significantly in terms of the amount of new bone and new blood vessels because the amount of rhBMP-2 was relatively low in this study. A limitation of the present study is that the release profile of BMP-2 from the scaffold has not been examined. Further studies are needed to determine the optimal doses of BMP-2 for bone formation within the scaffold.

[0168] The volume of vessels inside and outside the scaffold tended to be higher in the PG and PG+BMP-2 groups compared to the Empty group. There was no statistically significant difference in the amount of new bone between each group. And there tended to be more new bone in the PG and PG+BMP-2 groups than in the Empty group. The lack of significant difference was likely due to the small sample size. These results suggested the ability of the PG scaffold to promote neovascularization both inside and outside the scaffold, as well as to support new bone formation.

[0169] The reduction rate of scaffolds has been evaluated in previous reports by volume comparison using pCT [24a]. However, PCL’s low contrast makes it difficult to image in soft tissues. [24a]. In this study, PG could not be clearly visualized in the body. Therefore, tissue sections and preoperative pCT were compared instead. PCL has a slow degradation rate due to its hydrophobic nature. A previous study showed a volume change of 0.78% at 24 weeks after implantation of PCL alone in subcutaneous tissue. In addition to the polymer's chemical structure, factors such as surface modifications and the overall architecture of the scaffold play significant roles in its degradation process [3a]. It has been reported that mixing P-TCP and PCL increases hydrophilicity and degradation rates [24a]. Kawai et al reported that degradation rates of 20% TCP / 80% PCL scaffolds ranged from 10% to 25% at 8 weeks after implantation [25a] and 55% to 43% degradation when used combined with bone marrow-derived mononuclear cells [26a]. In this study, the reduction rate of the width of the PCL implants was 4.4% at the 8 weeks after surgery, corresponding to a volume reduction rate of 12.7%. This higher rate of change may be related to the fact that biodegradation is more pronounced in the bone defect area. The limitation is that we were not able to measure the reduction rate based on similar methods and are not able to follow the steric changes.

[0170] In this study, we used 3D-printed PG scaffolds to enhance the healing of large bone defects in beagle dogs. The implantation of PG scaffolds stimulated angiogenesis and osteogenesis both within and around the scaffold, demonstrating PG’s potential. However, adequate osteogenesis was not achieved by 8 weeks, indicating that PG alone or with minimal BMP-2 is insufficient for robust bone repair. This suggests that to achieve complete bone healing, higher concentrations of specific osteoinductive factors, such as BMP-2, or the combination of PCL with other supportive biomaterials may be required.[la] Samorezov JE and Alsberg E. Spatial regulation of controlled bioactive factor delivery for bone tissue engineering. Adv Drug Deliv Rev 2015; 84: 45-67. 20141129. DOI: 10.1016 / j.addr.2014.11.018.[2a] Habal MB and Reddi AH. Bone grafts and bone induction substitutes. Clin Plast Surg 1994; 21: 525-542.[3a] Siddiqui N, Asawa S, Birru B, et al. 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[0171] It is understood that the examples and embodiments described herein are for illustrative purposes only and that various modifications or changes in light thereof will be suggested to persons skilled in the art and are to be included within the spirit and purview of this application and scope of the appended claims. All publications, patents, and patent applications cited herein are hereby incorporated by reference in their entirety for all purposes.

Claims

WHAT IS CLAIMED IS:

1. A heat extrusion composition comprising polycaprolactone and a collagen substance, wherein the heat extrusion composition is a solid.

2. The heat extrusion composition of claim 1, wherein the collagen substance is distributed throughout the composition.

3. The heat extrusion composition of claim 1, wherein the collagen substance is selected from the group consisting of collagen, methacrylated collagen, gelatin, and methacrylated gelatin.

4. The heat extrusion composition of claim 1, wherein the collagen substance is not covalently attached to the polycaprolactone through a crosslinking moiety selected from an imine or an ether.

5. The heat extrusion composition of claim 1, wherein the collagen substance is present in nanosized particles.

6. The heat extrusion composition of claim 1, wherein solvent comprises less than 1% of the heat extrusion composition.

7. The heat extrusion composition of claim 1, further comprising an additive.

8. The heat extrusion composition of claim 7, wherein the additive is selected from the group consisting of a bioceramic additive, an anti-pathogenic additive, and an imaging additive.

9. The heat extrusion composition of claim 7, wherein the additive is selected from the group consisting of beta-tricalcium-phosphate, hydroxyapatite, calcium phosphate, biphasic calcium phosphate, aluminum oxide, zirconium dioxide, titanium oxide, zinc oxide, copper oxide, and magnesium oxide.

10. The heat extrusion composition of claim 7, wherein the additive is beta-tricalcium-phosphate.

11. The heat extrusion composition of claim 1, wherein the composition is a filament.

12. The heat extrusion composition of claim 1, produced by a process comprising:(a) contacting polycaprolactone, collagen substance, and an organic acid in a solvent;(b) casting the product of (a);(c) subjecting the product of (b) to evaporative conditions;(d) forming the product of (c) into pellets, thereby producing the additive manufacturing composition.

13. A method of making the heat extrusion composition according to claim 1, said method comprising:(a) contacting polycaprolactone, collagen substance, and an organic acid in a solvent;(b) casting the product of (a);(c) subjecting the product of (b) to evaporative conditions;(d) forming the product of (c) into pellets, thereby making the heat extrusion composition.

14. The method of claim 13, wherein the contacting comprises:(al) dissolving polycaprolactone and collagen substance in the solvent;(a2) adding the organic acid to the product of (al).

15. The method of claim 13, wherein the solvent is 1,1, 1,3,3, 3-hexafluoro- 2 -propanol (HFIP), and the organic acid is acetic acid.

16. The method of claim 13, wherein the ratio of polycarolactone to collagen substance is between 99:1 and 80:20.

17. The method of claim 13, wherein the (a) further comprises contacting the caprolactone, collagen substance, and organic acid with an additive.

18. The method of claim 17, wherein the additive is beta-tricalcium- phosphate.

19. The method of claim 13, wherein the (b) is cast into a low-form polytetrafluoroethylene evaporation dish at 70 °C for 8 hours under stirring.

20. The method of claim 13, wherein the solvent comprises less than 1% of the product of (c).

21. The method of claim 13, wherein the (d) is cutting the material sheet with stainless steel blades first laterally into 3mm strips and then orthogonally to form 3mm diameter rectangular pellets.

22. A heat extruded article, comprising two or more layers of the heat extrusion composition of claim 1.

23. The heat extruded article of claim 22, wherein the heat extruded article is formed by molten material extrusion.

24. A method of forming a heat extruded object comprising: heating and extruding the heat extrusion composition of claim 1 through a hot-melt extrusion nozzle for form an extrudate; and depositing the extrudate such that multiple layers are controllably deposited and fused forming the additive manufactured object.

25. The method of claim 24, wherein the extruding is by molten material extrusion.

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