Compositions and methods for preparation of biologic non-knee joint replacement

Abstract

The present disclosure provides compositions and methods for preparation of a novel patient-specific entirely biologic living non-knee joint arthroplasty / joint replacement.

Description

Docket 88800730-000525COMPOSITIONS AND METHODS FOR PREPARATION OF BIOLOGIC NON-KNEE JOINT REPLACEMENTCROSS REFERENCE TO RELATED APPLICATIONS

[0001] This invention claims the benefit of US Provisional Application Number 63 / 948,036 filed 23 December 2025; US Provisional Application Number 63 / 783,159 filed 03 April 2025; US Provisional Application Number 63 / 738,601 filed 24 December 2024; and US Provisional Application Number 63 / 738,587 filed 24 December 2024; each of which is incorporated by reference in its entirety.TECHNICAL FIELD

[0002] This disclosure relates to compositions and methods for the preparation of a biologic equivalent of a conventional non-knee replacement.BACKGROUND

[0003] Osteoarthritis (OA) is a highly prevalent and disabling condition that affects over 7% of people globally (528 million people). The prevalence of knee OA is rising due to the aging of the population and increasing rates of joint injury and, most importantly, obesity. Clinically, the field has focused on repairing local sites of cartilage damage to prevent spread of articular surface damage leading to a joint replacement. In situ repair for smaller focal lesions has relied primarily on marrow stimulation techniques which ultimately lead to tissue fill by inferior fibrocartilage representing a stop-gap solution at best. Emerging clinical strategies aim to use hydrogel polymers to augment the blood clot effect from marrow stimulation. For larger lesions, allografts or autologous osteochondral grafts are utilized, but are limited by tissue availability, and usually to a single compartment. Conventional tissue engineering strategies take several weeks to months to generate functional cartilage tissues. Similarly, clinically applied cell-based “tissue engineering” strategies (e.g., autologous chondrocyte implantation (ACI) and more emergent matrix-induced autologous chondrocyte implantation (MACI) technology)) require two surgeries and extensive rehabilitation (6-12 months) for return to activities such as running.

[0004] For joints afflicted with pervasive OA or in traumatic injury, such as, for example, a hip, a total joint replacement of metal and plastic is the clinical gold standard treatment option.Docket 88800730-000525However total hip replacement and other joint replacements have numerous complications and limitations associated with the current standards of care, including high infection and hardware failure rates, limited range of motion, persistent pain, nerve damage, and thrombosis. They also have a limited lifespans of 15 20 years, due to material failure or implant loosening / osteolysis, which means that younger recipients may need one or two revision implants. The demand for joint arthroplasties is projected to grow significantly from 2005 to 2030. Moreover, studies have found that cardiovascular disease is significantly increased in patients with OA compared with the general population, further amplifying its socioeconomic impact.

[0005] Therefore a biologic non-knee joint replacement design that can be used in lieu of standard metal and ultrahigh-molecular weight polyethylene (UHMWPE) artificial joints would represent a tremendous advance in the field.SUMMARY

[0006] This disclosure enables compositions and methods for preparation of a biologic non-knee joint for use in non-knee replacement. This disclosure at least partially addresses at least one of above inefficiencies. However, this disclosure can prove useful to other technical areas. Therefore, various claims recited below should not be construed as necessarily limited to addressing any of the above inefficiencies.

[0007] The present disclosure provides a biologic living non-knee joint replacement device, comprising: a first component comprising a first bone layer and a first cartilage layer; and a second component comprising a second bone layer and a second cartilage layer, wherein the first bone layer and the second bone layer each comprises a porous structure made of biocompatible and bioresorbable polymer, polymer blend, or polymer-ceramic composite and stem cell-derived bone cells, and wherein the first cartilage layer and the second cartilage layer each comprises a porous structure made of biocompatible and bioresorbable polymer or polymer blend and stem cell-derived cartilage cells. In certain embodiments, the device is produced by 3D printing technology. In some embodiments, the shape of the device is produced by scanning existing three part implants, combining the three parts into two components comprising the first component and the second component. In other embodiments, the device shape is modified using in silico finite element modeling to reduce stresses and improve surface conformity. In yet other embodiments, each of the first component and the second component comprises a cartilage domain, a boneDocket 88800730-000525domain, and an impermeable layer, and wherein the cartilage domain and the bone domain are separated by the impermeable layer. In still other embodiments, the first component further comprises an articular surface, wherein the second component further comprises an articular surface, wherein a cartilage subdomain is added to an articular surface of each of the first component and the second component to control lubrication properties. In additional embodiments, the mechanical properties of the device are designed to meet the mechanical demands of the joint under activities of daily living.

[0008] In further embodiments, the cartilage domain and the bone domain each include polymer blends, biomaterials, and biological factors to promote cartilage formation in the cartilage domain and bone formation in the bone domain and to promote phenotype maintenance after implantation. In still further embodiments, the stem cells are autologous stem cells. In other embodiments, the stem cells are autologous adipose-derived stem cells. In yet other embodiments, the stem cells are allogeneic stem cells. In still other embodiments, the stem cells are multipotent allogenic stem cells. In certain embodiments, the stem cells are human induced pluripotent stem cells.

[0009] In additional embodiments, the first cartilage layer and the second cartilage layer each comprises genipin crosslinked fibrin hydrogel. In further embodiments, the first cartilage layer and the second cartilage layer each comprises photocrosslinkable methacrylated hyaluronic acid (HAMA) hydrogel In some embodiments, the first bone layer and the second bone layer each comprises a polymer-ceramic composite. In other embodiments, at least one of the first cartilage layer or the second cartilage layer comprises a genipin-crosslinked fibrin (FibGen) hydrogel configured to delay degradation and provide improved compressive mechanical properties suitable for cartilage formation in vivo. In yet other embodiments, the first bone layer and the second bone layer each comprises a polymer-ceramic composite selected from polycaprolactone (PCL), polylactic acid (PLA), PLA / PBAT blends, or combinations thereof with ceramic particulates including hydroxyapatite (HA), P-tricalcium phosphate (P-TCP), zinc oxide (ZnO), or combinations thereof. In still other embodiments, the first cartilage layer and the second cartilage layer each comprises a polymer blend comprising PLA and PBAT configured to provide increased ductility and toughness relative to PLA alone.

[0010] In certain embodiments, at least one of the first component or the second component comprises a 3D-printed porous scaffold having an infill of about 70% to provide aDocket 88800730-000525porosity of about 30% with alternating raster orientations to balance pore size and strength. In additional embodiments, the scaffold pore architecture comprises an alternating infill pattern or a gyroid pattern, each configured to provide comparable equilibrium compressive modulus in the cartilage or bone domain. In further embodiments, the first articular surface and the second articular surface are each configured to provide increased congruence and reduced peak contact pressure relative to unmodified commercial geometries, as determined by finite element modeling of plastically deformed reference geometries. In some embodiments, the cartilage domain and the bone domain of at least one the first component or the second component are mechanically and spatially distinct from one another in a multi-layered construct configured to withstand peak loads representative of about ten times body weight without failure. In other embodiments, the stem cell derived bone cells and cartilage cells are obtained from autologous adipose-derived stem cells (ASCs) or from allogeneic human induced pluripotent stem cells (hiPSCs). In yet other embodiments, the cartilage layer comprises heparin-conjugated FibGen hydrogel configured to promote lubricin adsorption for boundary lubrication and reduced friction.

[0011] In additional embodiments, the bone layer includes ZnO nanoparticles embedded within a slow-degrading polymer fiber to provide slow zinc ion release and piezoelectric activity to promote osteogenesis. In certain embodiments, each of the first component and the second component comprises a 3D printed scaffold, each exhibiting pore sizes of about 200-400 pm and strut sizes of about 350-450 pm to support cellular infiltration and tissue integration. In further embodiments, the cartilage domain employs a ceramic-free polymer blend that is more compliant and ductile, and the bone domain employs a polymer-ceramic composite to assist in osteoinduction. In some embodiments, the cartilage layer hydrogel is deposited using a dualchannel nozzle enabling in situ mixing and deposition onto the scaffold surface with controlled thickness and infiltration depth by tuning feed rate and nozzle height. In other embodiments, the cartilage domain and the bone domain are selected to meet domain-specific mechanical targets derived from finite element modeling, including bone-domain tensile strength exceeding about 27 MPa and compressive strength exceeding about 63 MPa, and cartilage-domain tensile strength exceeding about 6 MPa and compressive strength exceeding about 48 MPa. In yet other embodiments, the device further comprises a multi-material scaffold, wherein the scaffold is additively manufactured using a multi-material printhead to produce a discrete cartilage domain and a discrete bone domain and an osteochondral architecture with non-planar articular contours.Docket 88800730-000525

[0012] In further embodiments, the cartilage layer and the bone layer are configured to interact without menisci. In yet further embodiments, the scaffold is designed to be entirely bioresorbable and non-immunogenic and chondro-inductive and osteo-inductive in vivo. In still further embodiments, the polymer-ceramic composite further comprises a biphasic ceramic of hydroxyapatite and P-tricalcium phosphate to promote osteoconduction and osteointegration. In certain embodiments, the device further comprises an articular layer geometry on at least one of the first component or the second component, wherein the articular layer geometry is defined using finite element deformation to target a desired average cartilage thickness and congruent contact under physiologic load.

[0013] The present disclosure also provides a method of preparing a biologic living nonkneejoint replacement device, comprising the steps of: a) fabricating a first component comprising a first bone layer and a first cartilage layer; and b) fabricating a second component comprising a second bone layer and a second cartilage layer. In certain embodiments, at least one of the first component or the second component is fabricated using 3D printing. In some embodiments, the first bone layer and the second bone layer each comprises a biocompatible and bioresorbable polymer-ceramic composite and stem cell derived bone cells and the first cartilage layer and the second cartilage layer each comprises a biocompatible and bioresorbable polymer blend and stem cell derived cartilage cells. In other embodiments, the polymer blend is a hydrogel and the cartilage cells and hydrogel are bioprinted into cartilage layers or deposited in the layer using a dip coating method via a negative mold. In additional embodiments, the method further comprises seeding the bone cells into the bone domain using a seeding device comprising a container configured to drive a flow of bone cells in media only into the bone regions, and preventing bone cells from reaching the cartilage region via an impermeable form fitting barrier.

[0014] The present disclosure additionally provides a biologic non-knee arthroplasty (TKA) / joint replacement device, comprising: at least one of a first component comprising a first bone layer and a first cartilage layer, or a second component comprising a second bone layer and a second cartilage layer, wherein the at least one of the first bone layer or the second bone layer comprises a biocompatible and bioresorbable polymer-ceramic composite and stem cell derived bone cells and the at least one of the first cartilage layer or the second cartilage layer comprises a biocompatible and bioresorbable polymer, polymer blend, or polymer composite and stem cell derived cartilage cells. In certain embodiments, the device is secured to a bone using ancillaryDocket 88800730-000525fixation bioresorbable pins, screws, or related fixation devices. In other embodiments, the device comprises a scaffold having four layers including: a coating layer to minimize friction; a cartilageseeding layer; an impermeable interface layer configured to reduce or prevent cell migration between domains; and a bone-seeding layer comprising a polymer-ceramic composite. In some embodiments, the coating layer and cartilage-seeding layer each comprise Ecovio T2308 with about 70% infill, and the bone-seeding layer comprises Ecovio T2308 with hydroxyapatite and 0-tricalcium phosphate. In yet other embodiments, the cartilage domain and the bone domain of the scaffold is optimized to a porosity of about 30%, with an infill density of about 70%, using sequential layers with rectilinear fibers alternating by about 45 degrees to enhance resistance to tensile and compressive failure. In still other embodiments, the cartilage domain comprises a triangular infill pattern to increase dynamic compressive modulus while maintaining interconnectivity for cellular communication. In certain embodiments of the device, finite element analysis of a post-in silico deformation geometry reduces peak contact pressure from about 27 MPa to about 14 MPa and yields more uniform contact stress distribution under near-full-extension running loads.

[0015] In some embodiments, the articular surface of the first component and the articular surface of the second component are prepared to receive heparin-conjugated hydrogel coating configured to promote lubricin adsorption for boundary lubrication upon exposure to synovial fluids. In other embodiments, the cartilage domain polymer blend comprises a PLA / PBAT blend selected to tune Young’s modulus and yield stress for the articular region. In further embodiments, the blend comprises between about 55 / 45 and about 75 / 25 PLA / PBAT by weight. In further embodiments, the cartilage domain polymer and hydrogel are printed using a modified dualprinting unit comprising a Y-shaped nozzle and stepper motor-controlled feed to achieve homogeneous coating thickness and controlled infiltration depth. In particular embodiments, the first component and the second component are configured for press-fit implantation with an added slab in the first component to accommodate the multilayer build. In additional embodiments, the device further comprises a custom guide to direct a first cut in a bone of a patient.

[0016] In certain embodiments, the bone domain polymer-ceramic composite comprises PCL with about 10 wt.% ZnO and about 20 wt.% HA / TCP, printed to achieve pore sizes of about 200-400 pm at about 6.5 mm by 6 mm by 1.2 mm dimensions for test coupons demonstrating cytocompatibility and ceramic loading fidelity. In some embodiments, the cartilage layer coatingDocket 88800730-000525is formed using hydrogel deposition at a nozzle height of between about 0.5 mm and about 1.0 mm above the scaffold surface to produce a uniform thin layer with controlled thickness by stepper motor speed. In some embodiments, domain-specific target properties are specified to ensure compressive yield stress exceeding about 25 MPa in both bone and cartilage regions and bonecartilage tensile ultimate strength exceeding about 10 MPa. In other embodiments, the scaffold features are configured to be manufactured with water-dissolvable support materials to preserve smooth articular surfaces after support removal. In additional embodiments, the ZnO component is included to provide piezoelectric activity to promote bone formation during physiologic loading without external electrodes.

[0017] In some embodiments, the osteochondral construct is dimensioned and contoured based on CAD models derived from micro-CT scans of conventional implants and modified by in silico plastic deformation to achieve improved congruence and reduced stress. In further embodiments, the cartilage layer and the bone layer are configured and validated by ex situ compressive loading up to about 10 body weights without gross damage and with loaddisplacement curves absent discontinuities indicating failure. In yet further embodiments, the cartilage layer polymer is subjected to ethylene oxide sterilization without significant change in tensile or compressive properties prior to implantation. In still further embodiments, an internal structure and a print fidelity of the first component and the second component are verified by micro-CT of triangular infill architectures assessing pore geometry, fiber width, pore height, and pore width. In particular embodiments, the device is compatible with ancillary bioabsorbable fixation devices used clinically for osteochondral allograft surgeries to provide initial compression and shear strength comparable to metallic fixation. In still other embodiments, the cartilage layer is engineered to recapitulate interstitial fluid pressurization and boundary lubrication mechanisms of native articular cartilage under sliding loads.

[0018] Additionally, the present disclosure provides a 3D printing nozzle, comprising: a first channel; and a second channel, wherein the nozzle is configured for printing one or more hydrogels, cells, and biomaterials onto a first scaffold layer via the first channel to support the growth of bone and onto a second scaffold layer via the second channel to support the growth of cartilage. In certain embodiments, the first channel is associated with a first hydrogel or biomaterial solution, and the second channel is associated with a second hydrogel or biomaterial solution. In some embodiments, the first hydrogel or biomaterial solution is a first blend ofDocket 88800730-000525hydrogel or biomaterials, and the second hydrogel or biomaterial solution is a second blend of hydrogels or biomaterials. In other embodiments, the first blend of hydrogel or biomaterials and the second blend of hydrogel or biomaterials are the same. In yet other embodiments, the first blend of hydrogel or biomaterials and the second blend of hydrogel or biomaterials are different. In still other embodiments, the first or second hydrogel or biomaterial solutions contain live cells. In further embodiments, the 3D printing nozzle is a screw-type 3D printing nozzle. In yet further embodiments, the 3D printing nozzle is a two-channel exit 3D printing nozzle. In additional embodiments, the one or more hydrogels is a photocrosslinkable hydrogel (UV HAMA).

[0019] The present disclosure further provides a modified dual 3D printing unit, comprising a 3D printing nozzle, comprising: a first channel; and a second channel, wherein the nozzle is configured for printing one or more hydrogels, cells, and biomaterials onto a first scaffold layer via the first channel to support the growth of bone and onto a second scaffold layer via the second channel to support the growth of cartilage.

[0020] The present disclosure additionally provides a 3D printer, comprising: a 3D printing nozzle comprising: a first channel; and a second channel, wherein the nozzle is configured for printing one or more hydrogels, cells, and biomaterials onto a first scaffold layer via the first channel to support the growth of bone and onto a second scaffold layer via the second channel to support the growth of cartilage, and a LAP photoinitiator, wherein the LAP photoinitiator is configured to activate the HAMA. In certain embodiments, the LAP photoinitiator is configured to emit UV light to activate HAMA crosslinking after deposition.

[0021] The present disclosure further provides an implant holder, comprising: a rotating platform configured to rotate a scaffold for a biologic living non-knee joint replacement device; and a securing element for securing the scaffold to the rotating platform, wherein the rotating platform is configured to rotate about 270 degrees to allow bioprinting across all curved surfaces of the scaffold.

[0022] The present disclosure also provides a method for bioprinting single-material and dual-formulation hydrogels to generate a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising: providing a bioprinter comprising a modified dual-channel nozzle configured to dispense hydrogel formulations under controlled pressure and feed rate; preparing a first hydrogel formulation comprising fibrinogen and thrombin crosslinked with genipin to form a fibrin-based hydrogel (FibGen); optionally preparing a secondDocket 88800730-000525hydrogel formulation comprising FibGen conjugated with heparin to promote lubricin adsorption for boundary lubrication; loading the first hydrogel formulation into a first channel of the nozzle and the second hydrogel formulation into a second channel of the nozzle; controlling dispensing parameters including feed rate and nozzle height to deposit the hydrogel formulations onto a porous polymer scaffold corresponding to an articular surface region of the component; selectively printing either: the first hydrogel formulation alone to form a single-material hydrogel coating; or both the first and second hydrogel formulations in a layered or blended configuration to form a dual -formulation hydrogel coating; infiltrating the hydrogel formulations into the scaffold pores to a predetermined depth to enhance mechanical integration and lubrication properties; and curing the printed hydrogel coating to achieve a target compressive modulus suitable for cartilage-like load-bearing performance.

[0023] Further, the present disclosure provides a method for bioprinting a single-material hydrogel coating on a three-dimensional printed scaffold, the method comprising: depositing a fibrin-based hydrogel onto a region of a three-dimensional printed scaffold corresponding to a cartilage-supporting domain; adjusting a nozzle height relative to the scaffold to form a thin hydrogel layer; and controlling a feed rate of the hydrogel during deposition to set a predetermined layer thickness, wherein the hydrogel comprises a genipin-crosslinked fibrin gel, optionally heparin-conjugated, and the scaffold is configured for a cartilage-supporting function. In certain embodiments, the method further comprises depositing the hydrogel using a printing nozzle at a height of about 0.5 mm to about 1.0 mm above the scaffold surface to obtain a uniform thin layer. In some embodiments, the method further comprises controlling the feed rate using a stepper motor, to set the thickness of the hydrogel layer during printing. In some embodiments, the hydrogel comprises a fibrinogen solution and a thrombin plus genipin solution at concentrations selected to achieve a desired compressive modulus under unconfined and confined compression.

[0024] The present disclosure additionally provides a method for bioprinting a dualformulation hydrogel on a three-dimensional printed scaffold, the method comprising: simultaneously dispensing a first hydrogel solution and a second hydrogel solution through a Y-shaped bioprinting nozzle having two channels that mix the solutions at or immediately prior to deposition; controlling a feed rate with a stepper motor to regulate hydrogel thickness and infiltration into the scaffold; and depositing the mixed hydrogel onto a scaffold surface corresponding to a cartilage-supporting region, wherein the Y-shaped nozzle comprises a two-Docket 88800730-000525channel exit configuration that yields homogeneous coating and infiltration. In certain embodiments, the first hydrogel solution comprises fibrinogen and the second hydrogel solution comprises thrombin and genipin, and further comprising mixing the first hydrogel solution and the second hydrogel solution immediately before exiting the nozzle to avoid conduit clotting. In some embodiments, the method further comprises tuning a nozzle height and a printing speed to achieve a homogeneous coating and controlled hydrogel infiltration into the scaffold. In other embodiments, the method further comprises applying a surface treatment to the scaffold to improve hydrogel coating homogeneity and infiltration.

[0025] The present disclosure also provides a method for layered bioprinting of hydrogels onto a multi-domain scaffold for osteochondral applications, the method comprising: providing a three-dimensional printed scaffold comprising a bone-supporting domain and a cartilagesupporting domain; depositing a hydrogel onto a first scaffold layer to support growth of bone; and depositing a hydrogel onto a second scaffold layer to support growth of cartilage, wherein the nozzle is configured to print hydrogels, cells, and biomaterials onto the respective scaffold layers. In some embodiments, the method further comprises dispensing a first hydrogel solution and a second hydrogel solutions through a bioprinting nozzle comprising a first channel associated with the first hydrogel solution and a second channel associated with the second hydrogel solution, and blending the first hydrogel solution and the second hydrogel solution at or before deposition. In other embodiments, the first and second hydrogel solutions are the same or different compositions. In additional embodiments, the hydrogel comprises heparin-conjugated genipin-crosslinked fibrin to promote lubricin adhesion and boundary lubrication. IN particular embodiments, the method further comprises adjusting a stepper-motor-controlled feed rate to adjust the thickness of the hydrogel layer. In certain embodiments, depositing the hydrogel is performed at room temperature or about 40°C. In other embodiments, deposition produces a hydrogel layer with compressive modulus increased by selecting genipin concentration in the hydrogel.

[0026] In yet other embodiments, a screw-type Y-nozzle is used to improve printability, and a two-channel exit nozzle is used to avoid bulge-up and to achieve superior coating outcomes. In still other embodiments, the scaffold comprises at least one of a three-dimensional printed polymer mesh configured for cartilage domain and bone domain, and further comprising infusing the polymer mesh with a cell-seeded hydrogel for chondroinductive and osteoinductive function; or a porous polymer phase and a ceramic phase selected to promote osteoconduction andDocket 88800730-000525osteointegration. In additional embodiments, the hydrogel is deposited onto a three-dimensional printed poly caprolactone or polylactic acid-based scaffold with pores in the range of 200-400 pm. In further embodiments, hydrogel printing parameters are selected to produce a homogeneous layer and controlled infiltration on three-dimensional printed polymeric scaffolds. In yet further embodiments, the printed hydrogel exhibits a compressive response characterized by zonedependent moduli under unconfined compression.

[0027] In still further embodiments, the hydrogel is bioprinted onto a NOVAJoint (nonknee) scaffold comprising four distinct printed layers: a rectilinear low-friction coating layer; a triangular cartilage-seeding layer; an impermeable rectilinear layer; and a triangular bone-seeding layer comprising Ecovio T2308 with hydroxyapatite and P-tricalcium phosphate. In certain embodiments, the method further comprises modifying the scaffold surface by chondroitin sulfate conjugation to reduce friction and improve lubrication. In other embodiments, the method further comprises coating the scaffold surface with polytetrafluoroethylene to reduce the friction coefficient of the articular surface. In some embodiments, deposition is performed with a multimaterial printhead to co-print materials corresponding to cartilage and bone regions of the scaffold. In yet other embodiments, extrusion temperatures for printing PLA and / or PBAT filaments are set at about 200°C to 220°C and the hydrogel is printed at room temperature or about 40°C. In further embodiments, feed rate and nozzle height are tuned such that the hydrogel thickness decreases with lower stepper motor steps per second and a nozzle height of about 1 mm yields a more even thin hydrogel layer relative to 0.5 mm. In still other embodiments, the method further comprises seeding cartilage cells in a methacrylated hyaluronic acid hydrogel with a photoinitiator and UV-curing to immobilize cells within cartilage regions of the scaffold.

[0028] In certain embodiments, the hydrogel formulation includes glycosaminoglycanmimetic polymers to maintain chondrogenic phenotype without significantly altering compressive moduli across deformation zones. In some embodiments, cartilage and bone domain compressive yield strengths are targeted to exceed 25 MPa with bone-cartilage interface tensile ultimate strength exceeding 10 MPa, as validated by mechanical testing. In other embodiments, porous scaffold infill density is set to about 70% infill with alternating fiber orientations to enhance resistance to failure in tension and compression. In yet other embodiments, print fidelity and fiber widths are validated by microCT and adjusted by extruder cleaning via cold-pull procedures to reduce microscopic defects. In still other embodiments, the bone domain scaffold furtherDocket 88800730-000525comprises a polymer-ceramic composite incorporating nano-hydroxyapatite, nano-P-tri calcium phosphate, and / or nano-zinc oxide to enhance osteoinductive properties. In further embodiments, simultaneous dispensing employs a modified dual printing unit equipped with a custom Y-shaped nozzle and stepper motor-controlled pressure to achieve homogeneous coating and controlled infiltration. In yet further embodiments, the hydrogel-coated scaffolds comprise a T2308 scaffold coated with FibGen and wherein the scaffolds exhibit low creep displacement under friction testing. In additional embodiments, in vitro culture of osteochondral constructs under reciprocal shear loading preserves cell viability and mechanical behavior over at least 28 days. In particular embodiments, cartilage hydrogel deposition and bone cell seeding are performed in chambers configured to position the first component and second component for targeted regional application. In certain embodiments, hydrogel deposition is coordinated with a sacrificial support structure printed for complex non-planar articular shapes and removed by water-dissolving support material to preserve surface smoothness. In some embodiments, the method further comprises optimizing the hydrogel-coated cartilage region to recapitulate interstitial fluid pressurization and boundary lubrication via lubricin adsorption. IN other embodiments, the method further comprises selecting and validating osteochondral scaffold porosity, fiber spacing, and infill multipliers to achieve target equilibrium and dynamic moduli for cartilage domain scaffolds. In still other embodiments, sterilization by ethylene oxide does not materially alter the compressive or tensile properties of the cartilage domain scaffold.

[0029] The present disclosure further provides a bioprinter ink nozzle for dispensing multicomponent mixtures to generate a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the nozzle comprising: (a) a proximal manifold configured to mount on a dual-syringe holder and receive first and second precursor streams respectively from a first syringe and a second syringe; (b) a bifurcated internal flow path including: (i) first and second inlet channels fluidically coupled to the proximal manifold; (ii) a Y-shaped junction having a bifurcation angle that merges the first and second inlet channels into a common outlet path to combine the precursor streams immediately prior to dispensing; and (iii) a terminal section comprising one of: (A) a screw-type static mixing segment configured to promote homogenous blending while limiting premature gelation; or (B) a two-channel co-exit segment configured to deliver laminar co-flow of the first and second streams without internal mixing, (c) a distal nozzle tip having an axisymmetric outlet geometry sized to deposit a continuous hydrogel filament withDocket 88800730-000525a target layer thickness and infiltration depth into an underlying porous scaffold; (d) a drive interface configured for microprecise control of dispensing by a stepping motor actuation of the syringes to set a feed rate independent of pneumatic pressure; (e) an adjustable standoff feature that sets a nozzle height relative to the scaffold surface to control the printed hydrogel layer thickness and infiltration; (f) a surface energy and wetting control treatment on at least the distal nozzle tip to improve filament continuity and reduce bulge formation during deposition; and (g) a modular tip architecture permitting interchangeable terminal segments to switch between the screw type static mixing segment and the two channel coexit segment, wherein the nozzle is configured to print single material formulations from the first syringe alone, dual formulation coflow from both syringes, or blended formulations produced within the Y -shaped junction or the screw type static mixing segment, so as to achieve homogenous hydrogel coating thickness and controlled infiltration depth on an articular surface region of the scaffold.

[0030] In certain embodiments, the bifurcation angle of the Y-shaped junction is between about 30° and about 90° to minimize residence time and premature gelation. In some embodiments, the screw-type static mixing segment comprises helical elements sized to maintain laminar flow at the dispensing feed rate established by the stepping motor. In other embodiments, the two-channel co-exit segment maintains separated laminar streams up to the outlet to enable boundary -layer deposition without bulk mixing. In further embodiments, the feed rate is adjustable in discrete step-per-second increments to achieve hydrogel layer thickness within a predetermined range. In yet other embodiments, the nozzle height is adjustable between about 0.5 mm and about 1.0 mm above the scaffold surface to tune infiltration into pores of about 200-300 pm. In additional embodiments, the distal tip includes a hydrophilic coating selected to reduce bulge-up during deposition of genipin-crosslinked fibrin hydrogel. In still other embodiments, the nozzle is mounted on a rotating platform configured to adjust angular orientation during printing to enable deposition of hydrogel formulations on complex non-planar scaffold geometries.

[0031] In yet further embodiments, at least one hydrogel formulation further comprises glycosaminoglycan (GAG) mimetics to promote chondroinduction and enhance cartilage-like extracellular matrix deposition within the scaffold. In still other embodiments, at least one hydrogel formulation or scaffold material incorporates nanoparticles selected from zinc oxide, hydroxyapatite, or P-tricalcium phosphate to mitigate immune response and promote glycosaminoglycan production and osteochondral integration. In certain embodiments, the nozzleDocket 88800730-000525is configured to bioprint at least one hydrogel formulation directly onto a composite scaffold comprising a porous polymer phase and a ceramic phase selected to promote osteoconduction and osteointegration, such that the hydrogel infiltrates the composite scaffold pores to enhance mechanical integration and chondroinductive performance.

[0032] The present disclosure additionally provides a bioprinting nozzle for dispensing a multi-component ink, comprising a first channel and a second channel that converge at a Y-shaped junction to deliver the components to an outlet, the nozzle being configured to print hydrogels, cells, and biomaterials onto a first scaffold layer and a second scaffold layer, wherein the first channel is associated with a first hydrogel or biomaterial solution and the second channel is associated with a second hydrogel or biomaterial solution. In certain embodiments, the nozzle geometry comprises a Y-shaped channel architecture selected from: a screw-type Y-nozzle including an internal mixing element, and a two-channel exit Y-nozzle in which the two channels meet at or immediately prior to the outlet. In some embodiments, the convergence location of the first and second channels is positioned to mix the first and second solutions at or immediately prior to discharge from the outlet to reduce or prevent in-conduit clotting when printing fibrinogen and thrombin plus genipin formulations. In particular embodiments, the nozzle further comprises a dispensing unit with two syringe holders operatively coupled to a stepper-motor-driven actuator configured to apply pressure and thereby micro-precisely control the extrusion feed rate of the multi-component ink. In yet other embodiments, at least one of the first hydrogel or biomaterial solution or the second hydrogel or biomaterial solution comprises live cells. In some embodiments, the first hydrogel or biomaterial solution and the second hydrogel or biomaterial solution comprise the same or different blends. In still other embodiments, the Y-shaped nozzle is selected from a screw-type configuration and a two-channel exit configuration as depicted in modified Y-shaped nozzle implementations. In additional embodiments, a modified dual printing unit includes the Y-shaped nozzle and a stepper motor control system to regulate feed rate and facilitate homogeneous hydrogel layer thickness and infiltration depth on a three-dimensional printed scaffold.

[0033] In certain embodiments, the two-channel exit Y-nozzle exhibits superior hydrogel coating outcomes relative to the screw-type Y-nozzle, which produces bulge-up structures in printed hydrogel. In other embodiments, the nozzle is integrated into a multi-material bioprinting setup used to deposit hydrogels onto designated cartilage regions of a NOVAIoint (non-knee)Docket 88800730-000525scaffold. In yet other embodiments, the dispensing unit replaces air-pressure actuation with a stepper-motor pressure drive to improve accuracy of printing rate for hydrogel deposition. In particular embodiments, the first and second channels are dedicated to first and second blends of hydrogel or biomaterials that may be the same or different, including formulations containing live cells.

[0034] The present disclosure also provides a modified dual bioprinting unit comprising a nozzle comprising a first channel and a second channel that converge at a Y-shaped junction to deliver the components to an outlet, the nozzle being configured to print hydrogels, cells, and biomaterials onto a first scaffold layer and a second scaffold layer, wherein the first channel is associated with a first hydrogel or biomaterial solution and the second channel is associated with a second hydrogel or biomaterial solution, and a motion stage, the unit being configured to achieve homogeneous hydrogel coating and controlled infiltration into three-dimensional printed polymeric scaffolds by coordinated control of feed rate and printhead movement.

[0035] Further, the present disclosure provides a method of bioprinting a multi-component hydrogel using a bifurcated nozzle, comprising: providing a first hydrogel solution and a second hydrogel solution to a Y-shaped nozzle having a first channel and a second channel; mixing the first and second solutions at or immediately prior to an outlet of the nozzle; dispensing the mixed solutions onto a scaffold surface while controlling the extrusion feed rate via a stepper motor and the relative motion of the printhead; and selecting a nozzle height above the scaffold to form a homogeneous thin hydrogel layer with controlled infiltration. In some embodiments, the nozzle is a two-channel exit Y-nozzle that yields superior hydrogel coating uniformity relative to a screwtype Y-nozzle that exhibits bulge-up in printed structures. In other embodiments, nozzle height is set between about 0.5 mm and 1 mm above the scaffold surface, and increasing the height from 0.5 mm to 1 mm produces a more even thin hydrogel layer. In certain embodiments, the feed rate of the hydrogel is controlled by the steps-per-second speed of the stepper motor to tune the printed layer thickness. In yet other embodiments, the first solution comprises fibrinogen and the second solution comprises thrombin and genipin, and in-conduit mixing prior to the outlet is avoided to prevent clotting within the nozzle. In further embodiments, the two solutions are mixed at or immediately before the outlet to prevent conduit clotting associated with T-shaped internal channel mixing of fibrinogen with thrombin plus genipin.Docket 88800730-000525

[0036] In other embodiments, the hydrogel layer thickness decreases with reduced stepper motor steps per second, demonstrating feed-rate dependent control of coating thickness. In particular embodiments, the method further comprises at least one of increasing the nozzle height from about 0.5 mm to about 1 mm above the scaffold surface to produce a more even thin hydrogel layer, or optimizing printing parameters including feeding rate, printhead speed, and scaffold surface treatment to achieve a controllable, homogeneous hydrogel coating and infiltration depth. In additional embodiments, the nozzle and printing parameters are configured to deposit heparin-conjugated genipin-crosslinked fibrin hydrogel onto a scaffold cartilage region while maintaining hydrogel mechanical performance. In yet other embodiments, micrographs of printed scaffolds confirm homogeneous hydrogel coating and infiltration depths between about 2 mm and 2.5 mm under optimized conditions. In still further embodiments, the nozzle and process parameters are tuned to yield homogeneous hydrogel layers across non-planar scaffold surfaces by coordinating nozzle height, stepper motor speed, and translation speed.

[0037] The present disclosure further provides a computer-implemented method for in silico modeling of plastic deformation to attain an articular surface shape for a scaffold of a biologic implant, the process comprising: obtaining scan data of a non-knee joint implant component comprising at least of a first component, a second tray and polyethylene insert by micro computed tomography (microCT), the scan data capturing articular and bone-interfacing geometries; constructing, from the scan data, surface models and a stereolithography (STL) file by segmenting and smoothing the scan data to reduce imaging artifacts and yield rendered surfaces; or segmenting and smoothing the scan data to construct rendered surfaces and a stereolithography (STL) file representing the articular and bone interfacing geometries; importing the STL file into a computer aided design (CAD) environment and generating CAD models of the first component and the second component, including modifying patient bone interfacing regions to define flat mating surfaces; generating a finite element analysis (FEA) mesh from the CAD models of the first component and the second component, the mesh defining distinct articular and bone scaffold domains; performing FEA with an elastic perfectly plastic material model applied to at least the articular layer, and applying a load representative of activities of daily living at near full extension, the load comprising a compressive force up to about ten body weights and, optionally, an internal / external rotational moment; inducing, in silico, plastic yielding over at least a portion of the articular layer to increase first-second congruence and reduce peak contact pressure, and savingDocket 88800730-000525a deformed geometry when the deformation approaches a target cartilage thickness; defining an articular surface shape based on the deformed geometry, including for the first component, extruding the deformed articular surface to form an articular domain, and for the second component, sectioning the CAD model at a desired average cartilage thickness and inclination and establishing tied contact across split domains; iteratively altering at least one of the second or the first articular geometry using results of the FEA to further reduce peak contact pressure below a yield stress threshold for a selected polymer blend, wherein one or more alterations result in a finalized articular surface shape; and exporting the finalized articular surface shape as manufacturing data for scaffold fabrication and subsequent mechanical validation.

[0038] In certain embodiments, the scan data is obtained by optical surface scanning after applying a flat, rapidly evaporating contrast paint, and the optical point cloud is converted to an STL file. In some embodiments, the plasticity is confined to the second articular layer and the first component remains elastic during the first deformation pass. In further embodiments, the load comprises approximately ten body weights at near full extension and includes an internal / external rotational moment to replicate range of motion. In additional embodiments, the finalized articular surface shape results in a peak contact pressure < about 14 MPa, which is below the yield stress of a PLA PB AT C8020 blend used for the articular layer. In yet other embodiments, the first articular surface is extruded to define an articular domain of target cartilage thickness prior to FEA validation.

[0039] The present disclosure further provides a computer-implemented method for generating a patient-specific articular surface shape for a non-knee implant by in silico plastic deformation, the method comprising: acquiring anatomical geometry of at least one component of a conventional non-knee implant via micro-computed tomography; segmenting scan datasets to generate surface models notwithstanding metal-induced artifacts; converting the segmented surface representations into a CAD model; generating a finite element analysis (FEA) mesh from the CAD model; applying boundary conditions and loads representative of activities of daily living to the FEA mesh and simulating elastic-perfectly plastic deformation of at least one articular component to increase articular surface congruence; saving the deformed geometry as an updated articular surface; and exporting the updated articular surface as a file suitable for manufacturing. In some embodiments, acquiring anatomical geometry further comprises rendering micro-CT scan data of metallic first and second components and a UHMWPE second insert, segmenting intoDocket 88800730-000525surface representations using Meshmixer, and reducing metal artifacts prior to CAD conversion. In other embodiments, converting to a CAD model comprises importing rendered surfaces into a CAD package and generating watertight CAD solids that are 3D-printable and suitable for downstream FEA. In further embodiments, simulating elastic-perfectly plastic deformation comprises modeling the second articular layer as elastic-perfectly plastic under approximately ten times body weight in near full extension to emulate worst-case running loads and saving a plastically deformed geometry that reduces peak contact pressures relative to the undeformed geometry. In yet other embodiments, the finite element loading protocol further comprises inducing plastic deformation at multiple kinematic states including internal / external rotation to replicate native range of motion and repeating the procedure allowing plastic deformation of the first component to enhance congruence across loading conditions.

[0040] In additional embodiments, the method further comprises generating a cartilage domain layer on a first component by applying a uniform pressure load to the articular surface in FEA to a target cartilage thickness, saving the deformed surface, and performing a surface extrusion to define an articular layer domain for subsequent mesh generation. In particular embodiments, the method further comprises generating a cartilage domain layer on a second component by splitting the CAD model with a sectioning plane at a desired average cartilage thickness and inclination and creating finite element meshes for the split domains with tied contact across the interface. In some embodiments, exporting the updated articular surface comprises writing the surface as an STL or CAD file suitable for use in slicer software for additive manufacturing and validation against micro-CT for print fidelity. In still other embodiments, the method further comprises validating the plastically deformed design by FEA under ten body weights at near full extension to confirm reduced peak contact pressures and more uniform contact areas relative to the pre-deformation geometry.

[0041] Additionally, the present disclosure provides a non-transitory computer-readable medium storing instructions that, when executed by one or more processors, cause a computer to perform a method for generating a patient-specific articular surface shape for a non-knee implant by in silico plastic deformation, the method comprising: acquiring anatomical geometry of at least one component of a conventional non-knee implant via micro-computed tomography; segmenting scan datasets to generate surface models notwithstanding metal-induced artifacts; converting the segmented surface representations into a CAD model; generating a finite element analysis (FEA)Docket 88800730-000525mesh from the CAD model; applying boundary conditions and loads representative of activities of daily living to the FEA mesh and simulating elastic-perfectly plastic deformation of at least one articular component to increase articular surface congruence; saving the deformed geometry as an updated articular surface; and exporting the updated articular surface as a file suitable for manufacturing. In some embodiments, the anatomical geometry acquisition further includes rendering and visualization using 3D Slicer prior to segmentation. In further embodiments, the method further comprises smoothing and modifying bone-contacting surfaces in the CAD model to create flat-mating interfaces to promote osteointegration of porous scaffolds. In certain embodiments, finite element material properties used for evaluating stress targets correspond to polylactic acid / polybutylene adipate terephthalate blends selected for the bone and cartilage domains. In yet other embodiments, the finite element analysis setup includes full-extension loading with peak forces approximating running, and contact analysis to compute contact pressure distributions across the first-second interface. In particular embodiments, the plastic deformation simulation yields a geometry in which the maximum contact pressure is less than the compressive yield stress of a prospective cartilage-domain polymer blend.

[0042] In some embodiments, the method further comprises fabricating physical prototypes of pre- and post-deformation designs by 3D printing and subjecting the prototypes to compressive loading up to ten body weights to verify absence of gross failure, with micro-CT scanning before and after loading. In other embodiments, mesh generation and solving are performed using a workflow that creates a finite element mesh from the SolidWorks-derived CAD model. In additional embodiments, the sectioning plane is oriented to impart a posterior inclination to the second cartilage layer prior to meshing and tied-contact definition. In certain embodiments, the method further comprises iterating plastic deformation steps across multiple flexion angles and internal / external moment conditions to converge on a geometry with improved congruence across the physiological range of motion. In particular embodiments, the exported articular surface shape is subsequently integrated into a multilayer scaffold design comprising a cartilage domain, an impermeable layer, and a bone domain for additive manufacturing.

[0043] In additional embodiments, the method further comprises documenting the in silico deformation workflow in a manufacturing process overview including initial scanning, CAD smoothing, in silico deformation to minimize peak stresses, and final design export for 3D printing. In certain embodiments, micro-CT image processing includes customized segmentation andDocket 88800730-000525smoothing algorithms to address imaging artifacts from metallic components prior to Meshmixer surface extraction. In other embodiments, CAD solids are generated in SolidWorks from rendered surfaces and validated by printing physical models prior to FEA. In particular embodiments, elastic-perfectly plastic behavior is applied to a second cartilage surrogate during FEA at approximately ten times body weight near full extension to plastically deform the second surface and increase congruence. In further embodiments, post-deformation FEA demonstrates a reduced peak contact pressure from about 27 MPa to about 14 MPa and a more uniform contact area relative to the initial commercial non-knee geometry. In some embodiments, generation of the first cartilage domain applies a uniform pressure to the articular surface to achieve a desired average thickness, and the deformed surface is extruded to create an articular layer domain suitable for further FEA. In other embodiments, generation of the second cartilage domain includes splitting with a sectioning plane at a desired average thickness and posterior inclination and meshing with tied contact to ensure kinematic continuity across layers. In yet other embodiments, the method further comprises ex situ mechanical verification by loading the 3D-printed first and second components to ten body weights in an Instron system at full extension, observing a smooth loaddisplacement curve without failure and confirming integrity by before / after micro-CT scans. In particular embodiments, the exported geometry is incorporated into a multilayer NOVAJoint (nonknee) scaffold comprising, in order, a low-friction rectilinear coating layer, a triangular cartilageseeding layer, an impermeable rectilinear layer, and a triangular bone-seeding layer. In other embodiments, the method further comprises validating print fidelity of porous regions using micro-CT analysis of fiber width and pore dimensions and implementing extruder cold-pull cleaning to reduce microscopic print defects.

[0044] Additionally, the present disclosure provides a method for designing and qualitycontrolling a scaffold for a biologic equivalent of a component of a conventional non-knee implant using a safety factor, the method comprising: generating a computer-aided design (CAD) model of a first component and a second component from scan data of a conventional non-knee implant, and defining in the CAD model an articular layer and a bone scaffold layer; constructing a finite element analysis (FEA) mesh of the first component and the second component and simulating a physiologic load case at near full extension comprising up to about ten body weights, to compute peak stresses in tension and compression for the articular layer and the bone scaffold layer; selecting one or more candidate flexible, degradable biopolymers or polymer blends, optionallyDocket 88800730-000525including polymer-ceramic composites for the bone scaffold layer, and determining material properties comprising at least tensile strength, compressive strength, and dynamic compressive modulus over a physiologic frequency range; applying a safety factor to the design for FEA-computed peak stresses to define minimum required properties for each layer, wherein the minimum required properties comprise at least: a required tensile strength > M x (peak tensile stress), and a required compressive strength > M * (peak compressive stress), for each of the articular layer and the bone scaffold layer; printing a scaffold with a predetermined porous architecture and infill pattern that balances load-bearing strength, hydrogel infiltration, and cellular communication, and controlling printing parameters to achieve dimensional fidelity; measuring the printed scaffold’s mechanical properties in tension and compression and its dynamic compressive modulus within a target range for the articular layer, and verifying that the measured properties meet or exceed the minimum required properties defined by the safety factor M; when the measured properties do not meet the minimum required properties, iteratively updating at least one of the scaffold geometry, infill percentage or pattern, layer composition, or polymer-ceramic content, repeating steps (b)-(f) until the measured properties meet the safety factor defined minimums; and releasing the scaffold for biological use upon verification that the safety factor-derived quality-control thresholds are satisfied for the articular layer and the bone scaffold layer.

[0045] In certain embodiments, the safety factor M is about 10. In other embodiments, the physiologic load case comprises an elastic perfectly plastic model applied to the articular layer during FEA. In some embodiments, the dynamic compressive modulus target for the articular layer is between about 15 MPa and about 60 MPa across 0.01-10 Hz. In further embodiments, the porous architecture comprises an alternating, triangular, honeycomb, or gyroid infill pattern with a porosity between about 25% and about 75%. In yet other embodiments, the bone scaffold layer comprises a polymer-ceramic composite containing nanoparticles selected from zinc oxide, hydroxyapatite, and 0 tricalcium phosphate. In particular embodiments, verification comprises preload microCT geometry capture and postload mechanical testing under compressive loads up to about ten body weights, followed by pass / fail classification against the safety factor-derived thresholds.

[0046] The present disclosure also provides a method of designing a multilayer osteochondral scaffold for a non-knee implant that satisfies a predefined safety factor, the method comprising: generating a finite element model of a first component and a second componentDocket 88800730-000525having separate bone and cartilage scaffold domains; determining peak tensile and compressive stresses in the bone and cartilage domains under a load representative of running at approximately ten times body weight; and specifying minimum tensile and compressive strengths for the bone and cartilage scaffold domains such that each minimum strength exceeds a corresponding peak stress by at least a safety factor of three. In some embodiments, the method further comprises selecting scaffold porosity and infill architecture to meet said minimum strengths, including choosing an infill density of about 70% to yield target yield strength and modulus while maintaining interconnected porosity for cell seeding. In other embodiments, the method further comprises validating the designed scaffold by fabricating test coupons and conducting tensile and compressive mechanical tests at physiologic or near-physiologic strain rates to confirm that measured strengths meet or exceed the specified minimum strengths corresponding to the safety factor. In certain embodiments, specifying minimum strengths comprises setting, for the bone domain, a minimum tensile strength greater than 27 MPa and a minimum compressive strength greater than 63 MPa, and, for the cartilage domain, a minimum tensile strength greater than 6 MPa and a minimum compressive strength greater than 48 MPa. In other embodiments, tensile and compressive testing of porous scaffolds printed with alternating triangular versus rectilinear patterns at 70% infill are used to confirm that the selected pattern meets or exceeds the safetyfactor-adjusted yield strength and toughness requirements. In additional embodiments, the method further comprises including risk-management activities under design control whereby hazards are identified and nonclinical testing (degradation, cytotoxicity, particulate formation, and implantation) is conducted, with redesign triggered if results are unacceptable. In other embodiments, acceptance criteria for scaffold mechanical properties are defined to sustain peak stresses predicted by finite element models with inclusion of a safety factor to minimize risk of failure.

[0047] The present disclosure further provides a quality-control method for production of a porous, additively manufactured osteochondral scaffold, comprising: printing a scaffold according to a selected infill pattern and nominal infill density; measuring actual porosity and fiber geometry by micro-computed tomography; comparing measured infill and geometry to nominal design values; and accepting or rejecting the scaffold based on whether the measured parameters fall within tolerances that ensure the scaffold will achieve the predefined safety -factor-adjusted minimum strengths under the finite element model load case. In some embodiments, the methodDocket 88800730-000525further comprises adjusting printer hardware by extruder “cold pull” cleaning to reduce microscopic print defects identified by micro-computed tomography and repeating the print-and-verify cycle until the acceptance criteria are satisfied. In other embodiments, micro-CT is used to validate infill density after printing by comparing mass / volume measurements to micro-CT-derived infill across multiple printed samples, and acceptance criteria require concordance sufficient to assure achievement of the safety-factor-based design strengths. In certain embodiments, the method further comprises implementing extruder cold-pull cleaning to reduce microscopic defects detected by micro-CT from approximately 15.4% of layers to at most about 1.51% of layers prior to lot release. In other embodiments, sterilization by ethylene oxide is verified not to alter tensile or compressive properties beyond acceptance limits for the cartilagedomain scaffold prior to lot release.

[0048] The present disclosure also provides a scaffold system comprising: a bone-domain lattice and a cartilage-domain lattice separated by an impermeable interface, each domain comprising a 3D-printed polymer or polymer blend configured such that the tensile and compressive strengths of each domain exceed finite-element-predicted peak stresses under a ten-times-body-weight load by at least a factor of three. In certain embodiments, the lattice design employs an alternating raster pattern at approximately 70% infill to balance pore size for cell seeding and mechanical integrity consistent with the safety factor requirements. In some embodiments, the scaffold comprises four layers including a 70% infill rectilinear coating layer, a 70% infill alternating triangular cartilage-seeding layer, a 100% infill impermeable layer, and a 70% infill alternating triangular bone-seeding layer with ceramic additives, each layer meeting safety-factor-adjusted strength targets.

[0049] Further, the present disclosure provides a method of verifying structural adequacy of an anatomically shaped non-knee scaffold, comprising: performing failure testing of additively manufactured first and second components at loads up to ten times body weight at full extension; comparing load-displacement curves and pre- / post-test micro-CT scans to finite element predictions; and confirming absence of gross damage consistent with the required safety factor margins. In particular embodiments, the materials comprise PLA-PBAT blends whose tensile and compressive properties are selected and confirmed to align with the safety-factor-based minimum strengths. In some embodiments, the finite element model uses PLA-PBAT C7575 for the bone domain and PLA-PBAT C6040 for the cartilage domain to represent design targets for strengthDocket 88800730-000525and modulus. In other embodiments, the running load case is modeled at near full extension and is used to size the safety-factor-adjusted minimum strengths by reducing peak contact pressure via articular surface congruence. In yet other embodiments, tensile dog-bone and compressive cylinder tests are performed to extract yield stress, ultimate stress, elastic modulus, and toughness for comparison to the predefined minimum strengths. In still other embodiments, print fidelity validation includes quantifying fiber spacing and width across layers and confirming stability of mass under early-stage degradation as part of release criteria. In some embodiments, the method further comprises cyclic durability testing under physiologic loading profiles and verifying continuous force-displacement curves without discontinuities over repeated loads, consistent with quality control criteria. In additional embodiments, the method further comprises cyclic durability testing under physiologic loading profiles and verifying continuous force-displacement curves without discontinuities over repeated loads, consistent with quality control criteria.

[0050] In further embodiments, the impermeable interface is a fully dense layer configured to prevent migration of cells between the cartilage and bone domains. In yet further embodiments, initial finite element models for design verification include a bone domain with 50% porosity PLA or Ecovio T2308 and a cartilage domain of PBAT, with peak domain stresses of approximately 9 MPa (tension) and 21 MPa (compression) for bone and 2 MPa (tension) and 16 MPa (compression) for cartilage, and the safety factor of three is applied to define minimum strengths. In other embodiments, the method further comprises documenting target properties for quality control including compressive yield stress greater than 25 MPa for both cartilage and bone domains and tensile ultimate strength greater than 10 MPa for the bone-cartilage interface.

[0051] The present disclosure also provides a method for assessing scaffold implant shape biofidelity upon fabrication using 3D printing and monitoring permanent deformation of a scaffold for a biologic equivalent of a component of a conventional non-knee implant or the implant itself, the method comprising: providing a scaffold comprising an articular surface region and a boneinterfacing region, the scaffold fabricated based on a computer-aided design (CAD) model derived from anatomic scan data of a conventional non-knee implant component; capturing a baseline three-dimensional geometry of the scaffold using a [laser] scanner to generate a point cloud representation of the articular surface region and the bone-interfacing region; converting the point cloud representation into a stereolithography (STL) file and comparing the STL file to the CAD model to determine a biofidelity metric comprising at least one of surface congruence, dimensionalDocket 88800730-000525tolerance, and curvature deviation; subjecting the scaffold to a mechanical load representative of physiologic conditions, the load comprising a compressive force up to about ten body weights applied at near full extension; rescanning the scaffold using the laser scanner after the mechanical load to generate a post-load point cloud representation; comparing the baseline and post-load point cloud representations to quantify permanent deformation of the articular surface region, the quantification comprising at least one of displacement magnitude, strain distribution, or change in curvature; and reporting the biofidelity metric and the permanent deformation quantification for validation of scaffold design prior to implantation.

[0052] The present disclosure further provides a non-destructive method of assessing biofidelity of an additively manufactured osteochondral scaffold relative to a nominal computer model, comprising: acquiring a three-dimensional scan of a printed scaffold; comparing measured architectural parameters from the scan, including at least fiber width, pore height, pore width and porosity, to nominal design parameters from a CAD or slicer model; and determining acceptance or rejection based on tolerances that ensure mechanical performance consistent with the design model. In certain embodiments, comparing measured architectural parameters further comprises calculating differences between nominal and actual fiber spacing across top, middle, and bottom layers to assess layer-wise fidelity. In some embodiments, the scan-derived porosity and infill density are used to explain measured differences in compressive yield strength and modulus between custom CAD patterns and slicer-generated patterns. In other embodiments, the scan modality is micro-computed tomography, and fiber width is quantified by layer to detect a significant decrease in top-layer fiber width relative to middle and bottom layers. In further embodiments, the method further comprises validating infill density by comparing mass / volume-derived infill with micro-CT-derived infill across multiple samples to establish micro-CT as a postprint fidelity assay. In yet other embodiments, scan-based architecture measurements include side pores, fiber width, pore height, and pore width in triangular infill scaffolds to assess biofidelity. In certain embodiments, scan-based measurements of infiltration depth for hydrogel coating on scaffolds are used to verify uniformity between approximately 2 and 2.5 mm across multiple locations. In particular embodiments, the three-dimensional scan comprises micro-computed tomography of porous scaffolds to quantify fiber spacing and width and to validate infill density against nominal design values. In still other embodiments, the three-dimensional scan is repeated after compressive or tensile loading of anatomical scaffold components, and pre- and post-testDocket 88800730-000525scans are compared to detect permanent deformation or gross damage. Tn yet further embodiments, the method further comprises quantifying porosity, strut size, and pore size from the scan to confirm consistency across polymer and polymer-ceramic compositions. IN additional embodiments, the scaffold comprises polymer-ceramic composites, and the scan confirms homogeneous ceramic distribution by surface analytical correlation. In some embodiments, scanbased porosity, strut size, and pore size are used to confirm consistency across polymer-ceramic compositions and to inform mechanical expectations under compression. In other embodiments, anatomical scaffold morphology is further verified in vivo by imaging post-explant to confirm minimal mass loss and no major shape change, consistent with pre- / post-scan assessments. In certain embodiments, acceptance tolerances are set such that the measured features and infill yield peak tensile and compressive stresses under a finite element load case that remain below material yield with a target safety factor.

[0053] The present disclosure also provides a non-destructive method of monitoring permanent deformation in an anatomical non-knee scaffold during failure testing, comprising: scanning first and second scaffold components before loading; loading the assembled components in compression to a target multiple of body weight at full extension; scanning the components after loading; and determining permanent deformation by image-based comparison of pre- and postload geometries and by correlation to load-displacement response. In some embodiments, the scanning modality comprises micro-computed tomography and the load-displacement response exhibits no sudden drop, indicating no gross failure consistent with the scan-based assessment. In other embodiments, permanent deformation monitoring is performed on patient-scale anatomical components fabricated at 70% infill and tested to approximately ten body weights. In further embodiments, the loading protocol includes static compressive loads up to approximately 3,000 N for anatomical implants with scan-based integrity checks before and after testing. In yet other embodiments, absence of permanent deformation is corroborated by smooth, continuous forcedisplacement curves over repeated loads and by pre- / post-scan comparisons without gross morphological change.

[0054] The present disclosure further provides a quality control method for printed scaffolds, comprising: performing non-destructive three-dimensional scanning to measure infill density and internal architecture; identifying microscopic print defects; implementing printer head cleaning; and reprinting until scan-derived defect rates and infill density meet acceptance criteria.Docket 88800730-000525In certain embodiments, identifying microscopic print defects comprises scan-based detection of intra-layer defects and validating that cleaning reduces detected defects to within acceptable limits. In some embodiments, printer head cleaning comprises an extruder “cold pull,” and scan-based defect analysis demonstrates reduction from about 15.4% of layers with defects to about 1.51% of layers.

[0055] The present disclosure additionally provides a system comprising: a scanner configured to acquire volumetric images of printed osteochondral scaffolds; and a processor configured to compute deviations in fiber width, pore height, pore width, strut size, porosity, and overall infill from nominal design values and to output pass / fail determinations for biofidelity and permanent deformation. In certain embodiments, the scanner is configured to quantify side pores, fiber width, pore height, and pore width in triangular infill scaffolds. In particular embodiments, the processor computes deviations and applies acceptance criteria aligned to target compressive yield stress, tensile strength at the cartilage-bone interface, and safety factor allowances.

[0056] The present disclosure also provides a method for scaffold failure analysis of a biologic equivalent of a component of a conventional non-knee implant, the method comprising: providing a scaffold comprising an articular surface region and a bone-interfacing region, the scaffold fabricated using 3D printing based on a computer-aided design (CAD) model derived from anatomic scan data of a conventional non-knee implant component; capturing a baseline three-dimensional geometry of the scaffold using micro-computed tomography (microCT) to generate volumetric imaging data of the scaffold structure; subjecting the scaffold to a mechanical load representative of physiologic conditions, the load comprising a compressive force up to about ten body weights applied at near full extension; rescanning the scaffold using microCT after the mechanical load to generate post-load volumetric imaging data; comparing the baseline and postload volumetric imaging data to identify structural changes indicative of failure, the structural changes comprising at least one of crack initiation, pore collapse, strut fracture, or permanent deformation of the articular surface region; and reporting the failure analysis results for validation of scaffold design and material selection. In some embodiments, the method, further comprises analyzing the micro CT scan to validate fabrication infill density and print quality.

[0057] Further, the present disclosure provides a non-destructive method for failure analysis of a biologic equivalent of a component of a conventional non-knee implant, comprising: acquiring pre-test micro-computed tomography scans of an additively manufactured first or secondDocket 88800730-000525component representing a biologic equivalent; loading the assembled components in compression at or near full extension to a target multiple of body weight; acquiring post-test micro-computed tomography scans of the components; and determining presence or absence of gross damage or permanent deformation by comparing the pre-test and post-test scans in view of the loaddisplacement response. In certain embodiments, acquiring the micro-computed tomography scans further comprises segmenting metallic or polymer components into surface representations, reducing metal-induced artifacts for conventional components, and importing the rendered surfaces into CAD for correlation to test geometry. In some embodiments, the loading step comprises compressive loading up to approximately ten body weights at full extension using a material testing system with custom jigs configured to hold first and second components in apposition. In other embodiments, assessing failure further comprises correlating micro-computed tomography-based findings with a smooth, continuous load-displacement curve to confirm absence of sudden drops indicative of failure. In yet other embodiments, the method further comprises validating print fidelity and internal architecture of the biologic equivalent prior to loading by quantifying porosity, strut size, and pore size from micro-computed tomography images. In still other embodiments, the method further comprises creating finite element models of the scanned components to predict peak stresses and contact pressures under a worst-case running load near full extension, and using the micro-computed tomography findings to confirm that test outcomes are consistent with finite element predictions. IN additional embodiments, the finite element models are generated from micro-computed tomography-derived renderings imported into CAD, and include separate articular and bone scaffold domains. In particular embodiments, the biologic equivalent comprises a 3D-printed porous scaffold fabricated at a defined infill percentage, and micro-computed tomography is used to verify actual infill prior to and after testing for consistency across specimens. In further embodiments, the presence of gross permanent deformation is determined by image-based comparison of pre- and post-load geometries and photographic inspection of components before and after testing. In yet additional embodiments, the method further comprising evaluating whether the post-test peak contact pressure or peak stress measured or inferred from the scans is consistent with a reduction in contact pressure achieved by in silico deformation of articular geometry relative to an undeformed design. In still further embodiments, the micro-computed tomography data are used to validate that the printed scaffold retains average pore size and strut size within specification after testing.Docket 88800730-000525

[0058] In certain embodiments, the micro-computed tomography scanning protocol is configured to quantify scaffold microarchitecture by layer, including side pores, fiber width, pore height, and pore width, for comparison to nominal design values. In some embodiments, the method further comprises validating infill density by comparing mass / volume-derived infill with micro-computed tomography-derived infill across multiple printed samples to establish microcomputed tomography as a post-print fidelity assay. In additional embodiments, scan-based defect analysis is used to monitor printing quality, and extruder cold-pull cleaning is implemented to reduce microscopic print defects from about 15.4% of layers to about 1.51% of layers prior to mechanical testing. In particular embodiments, the anatomical components are statically loaded up to approximately 3000 N in a material testing system at full extension as part of failure testing prior to cyclic or fatigue assessment. In other embodiments, absence of permanent deformation is corroborated by smooth, continuous force-displacement curves over repeated loads during ex situ testing of the assembled joint. In certain embodiments, micro-computed tomography measurements confirm that scaffold morphology and mass are maintained in vivo with minimal mass loss and no major shape change, consistent with pre- / post-scan assessments. In some embodiments, the method further comprises using finite element models to set target properties for compressive yield stress and interface tensile strength, and applying acceptance criteria during scan-based assessments aligned to those targets. In still other embodiments, scan-based architecture measurements further include verification of hydrogel infiltration depth into the scaffold within a target range across multiple locations. In further embodiments, the scan protocol and acceptance criteria are implemented within a design control framework for nonclinical testing, including degradation, cytotoxicity, particulate formation, and implantation assessments.

[0059] The present disclosure also provides a system for non-destructive failure analysis of a biologic equivalent of a conventional non-knee implant, comprising: a micro-computed tomography scanner configured to capture pre-load and post-load volumetric images of the components; and a processor configured to compare geometries and architectural measures to determine presence or absence of gross damage and to correlate the results with load-displacement testing data. In certain embodiments, the processor computes deviations in architectural metrics by layer, including detecting a significant decrease in top-layer fiber width relative to middle and bottom layers, to flag potential print-induced weakness prior to failure testing.Docket 88800730-000525

[0060] The present disclosure additionally provides a computer-implemented method for assessing biofidelity of a three-dimensional scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising: (a) acquiring imaging data of a printed scaffold, the imaging data comprising at least one of (i) micro-computed tomography (microCT) volumetric data; or (ii) imaging data suitable for geometrical measurement selected from optical imaging and scanning electron microscopy (SEM); (b) reconstructing a surface and volume representation of the scaffold from the imaging data, including: (i) segmenting and smoothing the imaging data to reduce artifacts; and (ii) generating a mesh or point-cloud model of the scaffold; (c) registering the reconstructed model to a nominal computer-aided design (CAD) model of the scaffold component to compute a shape biofidelity metric comprising at least one of root-meansquare (RMS) surface deviation, curvature deviation, or dimensional tolerance error; (d) extracting pore-structure metrics from the reconstructed model, comprising: (i) porosity percentage; (ii) pore size distribution; and (iii) strut thickness distribution; (e) computing spatial uniformity metrics of the pore structure across the scaffold, comprising at least one of nearest-neighbor statistics, regional variance, or anisotropy of infill orientation; (f) comparing the shape biofidelity metric and the pore-structure metrics to predetermined tolerances to determine consistency of the pore structure and overall scaffold biofidelity; and (g) generating a quality-control report indicating pass / fail and identifying regions of non-conformance for remedial manufacturing or processing.

[0061] The present disclosure further provides a computer-implemented method for assessing biofidelity of a three-dimensional printed scaffold that is a biologic equivalent of a component of a conventional non-knee implant, the method comprising: acquiring volumetric imaging data of the scaffold using micro-computed tomography or an equivalent three-dimensional scanning modality; segmenting the image data to reconstruct the scaffold geometry; extracting scaffold architectural metrics comprising at least fiber width, pore height, pore width, porosity, strut size, and overall infill density; comparing the extracted metrics to nominal design parameters derived from a computer-aided design or slicer model for the same scaffold; and generating a biofidelity score and pass / fail determination based on deviations from the nominal parameters within predefined tolerances. In certain embodiments, the method further comprising computing layer-wise architectural metrics by partitioning the volumetric data into top, middle, and bottom regions of the scaffold, calculating statistics for each region, and detecting layerdependent deviations that indicate print non-uniformity. In some embodiments, the comparingDocket 88800730-000525step comprises determining differences between nominal and measured infill density using both mass / volume-based estimations and micro-computed tomography-derived measurements, and reconciling discrepancies to validate print fidelity prior to use. In other embodiments, the method further comprising correlating the biofidelity deviations to predicted mechanical performance by mapping the deviations to finite element model peak tensile and compressive stresses for the scaffold and determining whether predicted stresses remain below material yield with a target safety factor. In yet other embodiments, the method further comprises detecting microscopic print defects from the volumetric imaging data, classifying the defects by layer occurrence, and outputting corrective actions to the printer process to reduce defect prevalence prior to acceptance.

[0062] In additional embodiments, the volumetric imaging data is also acquired before and after mechanical loading of anatomical scaffold components, and the method further comprises registering pre-load and post-load scans and quantifying permanent deformation or gross damage as an input to the biofidelity score. In some embodiments, the predefined tolerances for biofidelity acceptance are set to ensure that measured structural metrics across polymer and polymer-ceramic compositions remain within ranges shown to be consistently produced and mechanically acceptable. In particular embodiments, the method further comprises outputting a report that includes numerical deviations, layer-wise statistics, porosity maps, and a recommendation to accept, rework, or reject the scaffold. In some embodiments, extracting architectural metrics further comprises quantifying side pores, fiber width, pore height, and pore width in triangular infill scaffolds. In further embodiments, the biofidelity score weights discrepancies in porosity, strut size, and pore size according to their established consistency across polymer-ceramic formulations to maintain cross-composition comparability. In yet further embodiments, pre-load and post-load scanning is performed on patient-scale first and second components tested to approximately ten body weights at near full extension, and the absence of gross damage is used to update the biofidelity acceptance. In still further embodiments, the comparing step includes verifying that nominal infill settings selected for target pore size and strength, including approximately 70% infill, are achieved within tolerance by micro-computed tomography.

[0063] In some embodiments, the imaging modality comprises micro-computed tomography configured to quantify scaffold microarchitecture by layer, including side pores, fiber width, pore height, and pore width, for comparison to nominal design values. In other embodiments, validating infill density comprises comparing mass / volume-derived infill withDocket 88800730-000525microCT-derived infdl across multiple printed samples to establish microCT as a post-print fidelity assay. In certain embodiments, corrective actions include extruder cold-pull cleaning, and microCT defect analysis demonstrates reduction from about 15.4% of layers with defects to about 1.51% of layers. In additional embodiments, mechanical loading prior to post-load scanning comprises compressive loading of anatomical implants up to approximately 10 body weights with smooth, continuous load-displacement curves indicating no failure. In further embodiments, detecting layer-dependent deviations includes identifying a significant top-layer fiber width decrease relative to middle and bottom layers as a print-induced non-uniformity flag. In other embodiments, acceptance criteria for biofidelity incorporate target mechanical properties derived from finite element models, including compressive yield stress thresholds for cartilage and bone domains and tensile strength at the cartilage-bone interface. In still other embodiments, the method further comprises quantifying hydrogel infiltration depth into the scaffold from imaging data to verify uniformity between approximately 2 mm and 2.5 mm across multiple locations. In yet other embodiments, biofidelity tolerances and reporting are implemented within a designcontrol framework that includes risk identification and nonclinical testing to trigger redesign when results are unacceptable. In some embodiments, the method further comprises using biofidelity scan outcomes to corroborate ex situ or in vivo integrity assessments indicating minimal mass loss and no major change in scaffold morphology.

[0064] The present disclosure also provides a system comprising a volumetric scanner configured to acquire three-dimensional imaging data of printed scaffolds and a processor configured to perform a computer-implemented method for assessing biofidelity of a three-dimensional printed scaffold that is a biologic equivalent of a component of a conventional nonknee implant, the method comprising: acquiring volumetric imaging data of the scaffold using micro-computed tomography or an equivalent three-dimensional scanning modality; segmenting the image data to reconstruct the scaffold geometry; extracting scaffold architectural metrics comprising at least fiber width, pore height, pore width, porosity, strut size, and overall infill density; comparing the extracted metrics to nominal design parameters derived from a computer-aided design or slicer model for the same scaffold; and generating a biofidelity score and pass / fail determination based on deviations from the nominal parameters within predefined tolerances, and to output the biofidelity score, pass / fail status, and corrective action recommendations. In certain embodiments, the processor outputs pass / fail determinations aligned to validated print parametersDocket 88800730-000525for 70% infill and alternating triangular infill patterns selected to improve mechanical properties and resistance to failure.

[0065] Additionally, the present disclosure provides a non-transitory computer-readable medium storing instructions that, when executed by one or more processors, cause a computer to perform a computer-implemented method for assessing biofidelity of a three-dimensional printed scaffold that is a biologic equivalent of a component of a conventional non-knee implant, the method comprising: acquiring volumetric imaging data of the scaffold using micro-computed tomography or an equivalent three-dimensional scanning modality; segmenting the image data to reconstruct the scaffold geometry; extracting scaffold architectural metrics comprising at least fiber width, pore height, pore width, porosity, strut size, and overall infill density; comparing the extracted metrics to nominal design parameters derived from a computer-aided design or slicer model for the same scaffold; and generating a biofidelity score and pass / fail determination based on deviations from the nominal parameters within predefined tolerances. In certain embodiments, the processor executes image analysis pipelines to compute porosity, strut size, and pore size and to confirm that average pore size and strut size remain within specification.

[0066] The present disclosure also provides a computer-implemented method for assessing polymer damage in a three-dimensional scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising: (a) acquiring imaging data of a printed scaffold, the imaging data comprising at least one of: (i) micro-computed tomography (microCT) volumetric data; or (ii) imaging data suitable for geometrical and surface measurement selected from optical imaging and scanning electron microscopy (SEM); (b) reconstructing a surface and volume representation of the scaffold from the imaging data, including segmenting and smoothing the imaging data to reduce artifacts and generating a mesh or point-cloud model of the scaffold; (c) registering the reconstructed model to a nominal computer-aided design (CAD) model or to a baseline pre-load model of the scaffold; (d) computing one or more polymer-damage metrics from the registered models, the polymer-damage metrics comprising at least one of: (i) crack initiation indicators identified by discontinuities in strut connectivity or sharp local intensity / geometry changes; (ii) pore collapse or densification indicated by a decrease in regional porosity and loss of pore roundness; (iii) strut fracture or thinning identified by local reductions in strut thickness beyond tolerance bands; (iv) permanent deformation of the articular or bone-interfacing regions quantified by surface displacement magnitude or curvature change; or (v) changes in pore-sizeDocket 88800730-000525distribution or anisotropy of infill orientation indicative of damage propagation; (e) comparing the polymer-damage metrics to predetermined tolerances to classify damage severity into at least pass, rework, or fail; (f) generating a quality-control report highlighting regions of non-conformance and recommending remedial manufacturing or processing actions. In certain embodiments, the method further comprises suppressing artifacts from metal implants using MicroCT.

[0067] The present disclosure additionally provides a computer-implemented method for assessing polymer damage in a three-dimensional printed scaffold that is a biologic equivalent of a component of a conventional non-knee implant, the method comprising: acquiring volumetric imaging data of the scaffold using micro-computed tomography or an equivalent three-dimensional scanning modality; reconstructing the scaffold geometry from the imaging data; extracting architectural metrics comprising at least porosity, strut size, and pore size; computing layer-wise features comprising fiber width, pore height and pore width; comparing pre-defined nominal design parameters to measured values to determine deviations; and classifying polymer damage based on deviations and defect signatures, including detection of intra-layer microscopic defects and permanent deformation, and generating an accept / reject decision for the scaffold. In certain embodiments, the method further comprises acquiring pre-load and post-load volumetric scans, registering the scans, and quantifying permanent deformation and gross damage by comparing pre-load and post-load geometries together with load-displacement response from mechanical testing up to about ten body weights at or near full extension. In some embodiments, the comparing step further comprises validating infill density by reconciling microCT-derived infill with mass / volume-based estimates to confirm print fidelity prior to use in failure testing. In particular embodiments, the method further comprises mapping deviations in architectural metrics and measured defect prevalence to finite element model predictions of peak tensile and compressive stresses under worst-case loading to determine whether stresses remain below material yield with a specified safety factor. In other embodiments, the method further comprises detecting printer-induced microscopic defects from the volumetric imaging data, quantifying defect rates by layer, and outputting a corrective printing action comprising extruder cold-pull cleaning when the defect rate exceeds a threshold. In yet other embodiments, extracting architectural metrics further comprises quantifying side pores, fiber width, pore height, and pore width in triangular infill scaffolds to identify non-uniformities indicative of polymer damage or manufacturing defects.Docket 88800730-000525

[0068] In additional embodiments, permanent deformation is determined to be absent when pre-test and post-test microCT scans show no gross morphological change and the loaddisplacement response exhibits a smooth, continuous increase without sudden drops. IN certain embodiments, porosity, strut size, and pore size measurements are used to confirm that printing is consistent across polymer-ceramic composite compositions prior to mechanical testing for damage. In other embodiments, surface and bulk composition are verified by x-ray diffraction and attenuated total reflectance-Fourier transform infrared spectroscopy to exclude compositional changes that could confound damage assessment. In some embodiments, polymer damage classification further incorporates dynamic compressive moduli ranges to contextualize architectural deviations within cartilage-domain target properties. In yet other embodiments, mechanical loading replicates worst-case non-knee loads near full extension at approximately ten body weights and pre-scan / post-scan analysis is used to correlate contact pressure reductions achieved by in silico deformation with observed absence of damage. In further embodiments, polymer damage assessment also includes monitoring for early-stage degradation effects via imaging-correlated mass stability to ensure that observed features are attributable to damage rather than hydrolytic degradation. In still other embodiments, the imaging modality comprises microCT configured to quantify scaffold microarchitecture by layer, including side pores, fiber width, pore height, and pore width, for comparison to nominal design values. In particular embodiments, validating infill density comprises comparing mass / volume-derived infill with microCT-derived infill across multiple printed samples to establish microCT as a post-print fidelity assay. In additional embodiments, corrective actions include extruder cold-pull cleaning, and microCT defect analysis demonstrates reduction from about 15.4% of layers with defects to about 1.51% of layers.

[0069] In still other embodiments, the mechanical loading protocol includes compressive loading up to approximately ten body weights with smooth, continuous force-displacement curves and pre- / post-microCT scans indicating no gross damage. In certain embodiments, detecting layer-dependent deviations includes identifying a significant decrease in top-layer fiber width relative to middle and bottom layers as a print-induced non-uniformity flag. In further embodiments, acceptance criteria incorporate finite element model targets for compressive yield stress of cartilage and bone domains and tensile strength at the cartilage-bone interface. In other embodiments, the method additionally quantifies hydrogel infiltration depth from imaging data toDocket 88800730-000525verify uniformity between approximately 2 mm and approximately 2.5 mm across multiple locations. In some embodiments, scan-based polymer damage assessment is used alongside in vivo integrity findings indicating minimal mass loss and no major change in scaffold morphology post-implantation. In still further embodiments, the algorithm is implemented within a designcontrol framework that includes risk identification and nonclinical testing, with redesign triggered if results are determined to be unacceptable.

[0070] The present disclosure also provides a system comprising a volumetric scanner configured to acquire three-dimensional imaging data of printed scaffolds and a processor configured to perform a computer-implemented method for assessing polymer damage in a three-dimensional printed scaffold that is a biologic equivalent of a component of a conventional nonknee implant, the method comprising: acquiring volumetric imaging data of the scaffold using micro-computed tomography or an equivalent three-dimensional scanning modality; reconstructing the scaffold geometry from the imaging data; extracting architectural metrics comprising at least porosity, strut size, and pore size; computing layer-wise features comprising fiber width, pore height and pore width; comparing pre-defined nominal design parameters to measured values to determine deviations; and classifying polymer damage based on deviations and defect signatures, including detection of intra-layer microscopic defects and permanent deformation, and generating an accept / reject decision for the scaffold, and to output polymer damage metrics, a damage classification, and a pass / fail determination. In some embodiments, the processor is configured to compute porosity maps and layer-wise fiber width statistics to detect top-layer thinning relative to middle and bottom layers as an indicator of print-induced weakness. In other embodiments, the processor outputs pass / fail determinations aligned to validated print parameters for approximately 70% infill and alternating triangular infill patterns selected to improve mechanical properties and resistance to failure.

[0071] The present disclosure further provides a non-transitory computer-readable medium storing instructions that, when executed by one or more processors, cause a computer to perform a computer-implemented method for assessing polymer damage in a three-dimensional printed scaffold that is a biologic equivalent of a component of a conventional non-knee implant, the method comprising: acquiring volumetric imaging data of the scaffold using micro-computed tomography or an equivalent three-dimensional scanning modality; reconstructing the scaffold geometry from the imaging data; extracting architectural metrics comprising at least porosity, strutDocket 88800730-000525size, and pore size; computing layer-wise features comprising fiber width, pore height and pore width; comparing pre-defined nominal design parameters to measured values to determine deviations; and classifying polymer damage based on deviations and defect signatures, including detection of intra-layer microscopic defects and permanent deformation, and generating an accept / reject decision for the scaffold.

[0072] The present disclosure provides a method for controlled scaffold manufacturing and processing using flexible, degradable biopolymers for biological use to generate a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising: (a) selecting at least one flexible, degradable biopolymer from the group consisting of polylactic acid (PLA), polycaprolactone (PCL), polybutylene adipate terephthalate (PBAT), and blends thereof, optionally combined with ceramic fillers to promote osteoconduction; (b) preparing a printable filament or feedstock by melt blending the selected biopolymer or biopolymer blend under controlled temperature and shear conditions to achieve homogeneity and target filament diameter; (c) processing the blended biopolymer into a filament or pellet form suitable for fused deposition modeling (FDM) or direct bioprinting; (d) defining an appropriate multiplier for each polymer or feedstock to control amount of material extruded in 3D printing / FDM process using flexible biopolymers; (e) generating a computer-aided design (CAD) model of a scaffold corresponding to an articular surface region and a bone-interfacing region of the component, the scaffold comprising a porous architecture with a predetermined infill pattern and porosity to balance strength and cellular infiltration; (f) printing the scaffold using the prepared biopolymer feedstock under controlled extrusion parameters including nozzle temperature, layer height, infill percentage, and multiplier to achieve dimensional fidelity and mechanical integrity; (g) optionally incorporating a gradient in material composition or porosity between the articular region and the bone region to mimic native osteochondral structure; (h) post-processing the printed scaffold to remove support material and smooth the articular surface, and conditioning the scaffold under physiologic conditions to verify degradation rate and mechanical performance; and (i) validating the scaffold for biological use by confirming biocompatibility, degradability, and mechanical properties suitable for load-bearing in a non-knee joint. In certain embodiments, the bone layer’s polymer-ceramic composite comprises a biphasic bioceramic selected from hydroxyapatite (HA), - tricalcium phosphate (0 TCP), or a mixture of HA / 0 TCP.Docket 88800730-000525

[0073] In additional embodiments, the mixture of hydroxyapatite and P- tricalcium phosphate is present in a weight ratio of about 20 / 80 (HA / TCP ). In certain embodiments, the bioceramic is present as nanoparticles having an average particle size of approximately 100 nm to accelerate ion release and apatite formation. In some embodiments, the polymer of the bone layer comprises poly caprolactone (PCL) or PLA / PBAT and the composite further includes zinc oxide (ZnO) nanoparticles to promote osteogenesis and piezoelectric activity. In other embodiments, the PCL-ZnO or PLA / PBAT-ZnO composite further comprises the HA / p TCP mixture to provide dual osteoinductive and osteoconductive functionality. In yet other embodiments, the bone layer’s polymer-ceramic composite is configured to release calcium, phosphate, and zinc ions to upregulate vascular endothelial growth factor (VEGF) and support neovascularization. In still other embodiments, the bone layer includes a spatial gradient of bioceramic content that increases toward the bone interfacing region to enhance osteointegration while maintaining toughness. In additional embodiments, the polymer-ceramic composite is verified by thermogravimetric analysis (TGA) to confirm post print bioceramic loading within ±10% of target weight percent. In further embodiments, the polymer-ceramic composite is characterized by x ray diffraction (XRD) and attenuated total reflectance-Fourier transform infrared spectroscopy (ATR FTIR) to confirm phase and chemical integrity after processing. In yet further embodiments, energy dispersive X-ray analysis (EDXA) mapping demonstrates the surface distribution of ceramic nanoparticles on struts of the porous architecture. In particular embodiments, the porous architecture in the bone layer comprises pores of about 200-400 pm and strut sizes of about 300-350 pm to balance osteoconduction and vascular ingress. In some embodiments, the polymer-ceramic composite exhibits a porosity selected from about 25% to about 75% and an infill pattern selected from alternating, triangular, honeycomb, or gyroid to tune load-bearing and microvascular perfusion. In certain embodiments, the bone layer’s polymer-ceramic composite maintains dimensional fidelity and consistent bioceramic distribution after printing, as evidenced by optical imaging and scanning electron microscopy. In further embodiments, the inclusion of HA / p TCP and ZnO is selected to achieve early osteoid deposition and sustained angiogenic signaling for at least the first 4-8 weeks post-implantation .

[0074] The present disclosure also provides a method for creating a ceramic containing layered biologic, comprising: solvent casting of ceramic nanoparticles evenly into solubilized polymer using mixing or sonication between about 10% to about 40% w / v; evaporation of excessDocket 88800730-000525solvent and creation of polymer-ceramic composite flakes; shredding and granulation of the polymer-ceramic composite flakes; evaluating the shredded and granulated polymer-ceramic composite for melt flow index for flowability in melt blending; melt blending of the polymerceramic composite and extrusion of polymer-ceramic into filament for use in 3D printing and; 3D printing, fused deposition modeling, or direct printing of polymer-ceramic composite feedstock into controlled layers, scaffolds, and structures.

[0075] The present disclosure further provides a method for controlled manufacturing and processing of a three-dimensional printed scaffold formed of a flexible, degradable biopolymer for biological use as a biologic equivalent of a component of a conventional non-knee implant, the method comprising: selecting a biocompatible and bioresorbable polymer or polymer blend comprising polylactic acid (PLA), polycaprolactone (PCL), polybutylene adipate terephthalate (PBAT), or a blend thereof; fabricating scaffold fdaments by melt blending and extrusion to obtain printable filaments; three-dimensionally printing a porous scaffold having a predetermined pore architecture and infill density configured to provide targeted mechanical properties; and validating that the scaffold’ s tensile and compressive properties meet predefined strength and modulus targets suitable for non-knee loads. In some embodiments, the polymer blend comprises a PLA-PBAT composition extruded into filaments and spooled for printing, and the scaffold is printed at an infill density of about 70% with a raster architecture selected from alternating or triangular patterns to balance pore size and strength. In certain embodiments, the method further comprises controlling printer parameters including nozzle size, layer height, extrusion width and thickness, and print speed to achieve target strut cross-sectional area and pore dimensions, thereby tuning yield strength, modulus, and toughness under physiologic strain rates. In other embodiments, the method further comprises conducting non-destructive micro-computed tomography analysis to quantify porosity, strut size, pore size, fiber spacing, and infill density pre- and post-processing, and accepting or rejecting based on conformance to nominal CAD-derived specifications. In yet other embodiments, the method further comprises process qualification steps including: verifying crystallinity and thermal response of printed blends; verifying that solvent-based or melt processing does not change composition by ATR-FTIR and XRD; and confirming printed composite content by thermogravimetric analysis.

[0076] In particular embodiments, the scaffold comprises a cartilage-domain polymer blend selected for ductility and dynamic compressive moduli within targeted cartilage ranges, andDocket 88800730-000525a bone-domain composite including ceramics selected for osteoinduction while maintaining toughness and reduced brittleness via ductile polymer matrix. In additional embodiments, the method further comprises a quality-control loop in which micro-CT-detected microscopic print defects are used to trigger corrective printer maintenance comprising extruder cleaning and reprinting until defect rates and infill accuracy achieve acceptance criteria. In some embodiments, the method further comprises degradation assessment under physiologic aqueous conditions to confirm mass stability and maintenance of compressive properties over a defined interval, thereby qualifying the biopolymer scaffold for biological use. In other embodiments, the method further comprising ex situ failure testing of anatomically shaped first and second components at loads up to about ten times body weight, with micro-CT pre- and post-tests and load-displacement analysis to confirm absence of gross damage consistent with design targets. In further embodiments, the PLA-PBAT blends comprise custom ratios including 60 / 40 and 75 / 25 PLA / PBAT, and properties evolve monotonically with composition for yield stress, modulus, and toughness. In yet further embodiments, the selected infill pattern at about 70% infill provides an optimal pore size on the order of approximately 0.3 mm and superior strength compared to lower infill densities.

[0077] In certain embodiments, the method further comprises selecting triangular pore geometry to improve compressive equilibrium and dynamic moduli relative to rectilinear and alternating patterns. In some embodiments, increasing strut cross-sectional area via extrusion width and thickness control increases elastic modulus and yield strength of PLA prints while maintaining toughness. In other embodiments, ATR-FTIR and XRD confirm no unanticipated compositional or structural changes arising from solvent casting or printing of PLA / PBAT- or PCL-ceramic composites. In yet other embodiments, cartilage-domain porous polymers exhibit dynamic compressive moduli within a target range, with ranking PLA greater than PCL greater than PBAT across matched porosity. In still other embodiments, bone-domain polymer-ceramic composites comprise PLA / PBAT or PCL with nano-hydroxyapatite, nano- -TCP, and / or nano-ZnO to promote osteogenesis and integration while avoiding excessive brittleness. In further embodiments, micro-CT-derived porosity, strut size, and pore size confirm consistent printing results across multiple polymer-ceramic compositions. In yet further embodiments, melt blending and extrusion are performed either by two-stage melt blending and re-extrusion or by solvent casting followed by shredding and extrusion into filaments suitable for printing. In still furtherDocket 88800730-000525embodiments, sterile processing and testing demonstrate that ethylene oxide sterilization does not change tensile or compressive mechanical properties for cartilage-domain scaffolds.

[0078] In additional embodiments, the flexible degradable biopolymers comprise PBAT and PLA blends with documented elastomeric behavior and toughness improvement in PLA by melt blending with PBAT, with extrusion into fdaments for scaffold printing. In certain embodiments, the internal structure is optimized at approximately 70% infill density with alternating rectilinear orientations to provide enhanced resistance to failure in tension and compression. In particular embodiments, process parameters for printing PLA and / or PBAT include extrusion temperatures of about 200-220°C for scaffolds, and hydrogel printing at room temperature or about 40°C, to maintain print fidelity and compatibility with biological use. In some embodiments, micro-CT is used to quantify side pores, fiber width, pore height, and pore width for triangular infill scaffolds to assess microarchitecture against nominal design. In other embodiments, micro-CT defect analysis of porous samples is used to monitor printing quality before and after extruder cold-pull cleaning, demonstrating defect reduction from about 15.4% of layers to about 1.51% of layers. In yet other embodiments, in vitro hydrolytic degradation testing of cartilage and bone scaffold compositions shows minimal mass loss over months with maintained morphology and mechanical integrity, including faster degradation for ceramic-containing composites relative to polymer alone. In further embodiments, fatigue durability of Ecovio T2308 scaffolds at 70% alternating triangular infdl meets at least one million cycles at 4 MPa and 3000 RPM, with most samples achieving up to five million cycles without failure. In yet further embodiments, acellular scaffold properties are maintained after ethylene oxide sterilization, with no change in tensile or compressive mechanical properties for cartilage-domain materials. In still further embodiments, bone-domain composites comprising Ecovio T2308 with 30 wt.% hydroxyapatite / p-TCP exhibit sufficient tensile and compressive properties and promote cell adhesion and osteoinduction. In particular embodiments, ex situ mechanical testing of anatomically scaled NOVAJoint (non-knee) constructs demonstrates sustained integrity under applied loads with continuous force-displacement curves over repeated loading cycles during culture.

[0079] The present disclosure additionally provides a system for controlled manufacturing and processing of a flexible, degradable biopolymer scaffold for a biologic equivalent non-knee implant component, comprising: a filament production line configured for melt blending PLA andDocket 88800730-000525PBAT and spooling filaments; a fused deposition modeling printer configured to generate user-defined porous architectures at target infill; and a quality -control station comprising micro-CT and spectroscopic analysis to verify architecture, composition, and mechanical suitability for non-knee loading.

[0080] The present disclosure further provides a layered biologic non-knee implant, comprising: (a) an anatomically shaped scaffold body printed in a single, continuous three-dimensional (3D) build without post-print assembly seams, the scaffold body defined by a computer-aided design (CAD) model derived from scan data of a conventional non-knee implant component; (b) a cartilage layer forming an articular surface region of the scaffold body, the cartilage layer comprising a flexible, degradable polymer or polymer blend and a hydrogel coating comprising a genipin-crosslinked fibrin hydrogel configured for boundary lubrication and interstitial fluid pressurization; (c) a bone layer forming a bone-interfacing region of the scaffold body, the bone layer comprising a polymer-ceramic composite selected to promote osteoconduction and osteointegration; (d) a printed boundary layer that is co-formed during the single, continuous 3D print to transition between the cartilage layer and the bone layer, the boundary layer comprising a gradation in at least one of material composition, porosity, or infill orientation to maintain mechanical continuity and enhance cellular communication across layers; (e) an internal porous architecture within at least one of the cartilage layer, boundary layer or bone layer, having a predetermined infill pattern and porosity selected to balance load-bearing strength and hydrogel infiltration depth; and (f) a surface-smoothing feature achieved by printing with a water-soluble support material that is removed post-print to reduce micro-abrasions on the articular surface region, wherein the anatomically shaped scaffold body is configured to reduce peak contact pressure under physiologic loading and to provide layer-specific mechanical properties suitable for cartilage-like tribology at the articular surface and for osteoconductive integration at the boneinterfacing region. In certain embodiments, the scan data comprises micro-computed tomography (microCT) of the conventional non-knee implant component. In some embodiments, the flexible, degradable polymer or polymer blend of the cartilage layer comprises a PLA-PBAT blend tuned for ductility and toughness.

[0081] In other embodiments, the hydrogel coating comprises heparin-conjugated HAMA or FibGen to promote lubricin adsorption for boundary lubrication. In additional embodiments, the polymer-ceramic composite of the bone layer comprises PLA / PBAT or PCL with nanoparticlesDocket 88800730-000525selected from zinc oxide, hydroxyapatite, and P-tricalcium phosphate. In particular embodiments, the predetermined infdl pattern is selected from alternating, triangular, honeycomb, or gyroid, with a porosity between about 25% and about 75%. In further embodiments, the boundary layer includes a graded change in infill orientation (0°, 45°, 90°, 135°) to enhance mechanical continuity across the cartilage and bone layers. In certain embodiments, printing is performed with a multimaterial printhead to deposit distinct materials for the cartilage layer, boundary layer, and bone layer in a single, continuous build. In some embodiments, the anatomically shaped scaffold body is derived from scan-to-STL-to-CAD processing that introduces flat mating surfaces at bone interfaces while preserving articular congruence. In other embodiments, the water-soluble support material comprises polyvinyl alcohol (PVA) to maintain articular surface smoothness upon removal.

[0082] Additionally, the present disclosure provides a layered, multifunctional biologic non-knee implant, comprising: a first component having a cartilage-domain lattice and a bone domain lattice; and a second component having a cartilage domain lattice and a bone domain lattice, wherein each domain comprises a biocompatible, bioresorbable polymer or polymer blend, or a polymer ceramic composite appropriate for its biological function, and the implant is configured for patient-specific anatomy. In certain embodiments, the cartilage domain lattice and the bone domain lattice are provided as separate domains with an interface boundary between them, and the domains are designed to withstand physiologic non-knee loads up to about ten body weights under finite element based design criteria. In some embodiments, the cartilage domain uses a more compliant, ductile polymer blend and the bone domain uses a polymer ceramic composite to promote osteoinduction and osteointegration while maintaining toughness. In other embodiments, the layered construct comprises four distinct layers including: a low-friction rectilinear coating layer for the articular surface; a 70% infill alternating triangular cartilage-seeding layer; an impermeable rectilinear layer to prevent cell migration; and a 70% infill alternating triangular bone-seeding layer containing hydroxyapatite and P-tricalcium phosphate.

[0083] In further embodiments, CAD renderings show sliced regions of the layered first and second components with the layered configuration implemented throughout the anatomical geometry. In additional embodiments, the internal structure is optimized at about 70% infill, with rectilinear fiber orientations alternating by 45° in consecutive layers to enhance resistance to failure in tension and compression. In some embodiments, triangular pore geometry is selected inDocket 88800730-000525at least one domain to improve compressive equilibrium and dynamic moduli, with alternating or stacked triangular layers selected to improve cellular communication while maintaining modulus. In further embodiments, layer-specific polymers include Ecovio T2308 for at least the cartilage-associated layers and Ecovio T2308 with hydroxyapatite / p-TCP for the bone-associated layers. In particular embodiments, design acceptance criteria are tied to finite-element model-based targets including compressive yield stress thresholds for cartilage and bone regions and tensile strength of the cartilage-bone interface. In some embodiments, the layered construct is validated in ex situ loading up to approximately ten body weights on an Instron system with continuous force-displacement curves and no gross damage on pre / post scans prior to cadaver testing. In other embodiments, in vivo integrity assessments of multilayer constructs demonstrate minimal mass loss and no major change in scaffold morphology while maintaining compressive yield strength compatible with design criteria. In further embodiments, the cartilage-domain polymer is a more compliant, ductile blend and the bone-domain composite includes bioceramics and optionally nano-zinc oxide to promote osteogenesis while mitigating brittleness via the ductile matrix.

[0084] The present disclosure also provides a method of manufacturing a layered, multifunctional biologic non-knee implant, comprising: importing patient specific surface models into CAD; generating a layered scaffold design having discrete cartilage and bone domains; and 3D printing the layered scaffold in a single, continuous build using a multi material printhead to assign different materials to the respective cartilage and bone regions. In some embodiments, the method further comprises printing complex, non planar anatomical shapes with sacrificial supports and removing the supports with a water soluble material to maintain smooth articular surfaces of the printed anatomy. In certain embodiments, the multi-material print process uses a printer upgraded with a multi-material printhead to simultaneously assign and continuously print different materials for cartilage and bone regions in one job. In particular embodiments, printing of anatomical non-planar shapes uses water-dissolvable PVA support to avoid micro-abrasions on articular surfaces caused by mechanical support removal. In additional embodiments, extrusion temperatures for PLA / PBAT materials are maintained about 200-220°C, with separate hydrogel deposition at room temperature or approximately 40°C to maintain fidelity and compatibility with biologic payloads. In other embodiments, the method further comprises micro-CT analysis to quantify side pores, fiber width, pore height, and pore width of a triangular-infill scaffold to assessDocket 88800730-000525fidelity to nominal design by layer. In yet other embodiments, micro-CT-based print fidelity validation is corroborated by mass / volume infill calculations across multiple printed samples, establishing micro-CT as a post-print infill assay. In still other embodiments, the method further comprises a printer maintenance step in which an extruder cold-pull cleaning protocol is executed to reduce micro-CT-detected layer defects from about 15.4% of layers to about 1.51% of layers prior to lot release.

[0085] The present disclosure further provides a method for preparing a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising: (a) providing a three-dimensional (3D) printed scaffold comprising at least one of a first component or a second component, the scaffold including a bone-interfacing region and an articular surface region, the scaffold formed from a biocompatible and bioresorbable polymer or polymer-ceramic composite; (b) subjecting at least a portion of the scaffold to plasma treatment under controlled pressure and exposure time to modify surface energy and enhance cell adhesion and hydrogel infiltration; (c) rinsing or conditioning the plasma-treated scaffold in ethanol (EtOH) to remove residual contaminants and to sterilize the scaffold prior to biological use; (d) optionally repeating the plasma treatment and ethanol conditioning steps to achieve a predetermined surface wettability and sterility level; (e) sterilizing the scaffold using ethylene oxide; (f) drying the scaffold under aseptic conditions and verifying that the scaffold maintains dimensional fidelity and porosity after plasma and ethanol processing; and (g) releasing the scaffold for subsequent cell seeding or hydrogel infusion upon confirmation of sterility and surface activation. In certain embodiments, the plasma treatment comprises oxygen plasma applied at a pressure between about 0.1 Torr and about 1.0 Torr for a duration between about 30 seconds and about 10 minutes. In some embodiments, the ethanol conditioning comprises immersion in 70% EtOH for a time period between about 5 minutes and about 30 minutes. In other embodiments, the scaffold comprises a porous architecture with an infill pattern selected from alternating, triangular, honeycomb, or gyroid, and a porosity between about 25% and about 75%. In further embodiments, the scaffold comprises a polymer-ceramic composite including PLA / PBAT or polycaprolactone (PCL) and nanoparticles selected from zinc oxide, hydroxyapatite, and P-tricalcium phosphate. In yet other embodiments, the plasma treatment and ethanol conditioning steps are performed sequentially in a closed sterile chamber to maintain aseptic conditions.Docket 88800730-000525

[0086] The present disclosure provides a method of preparing a three-dimensional printed scaffold for biological use as a biologic equivalent of a component of a conventional non-knee implant, comprising: fabricating a porous scaffold from biocompatible and bioresorbable polymers, polymer blends, or polymer-ceramic composites by fused deposition modeling to achieve a target pore architecture and infdl suitable for non-knee loading; characterizing the printed scaffold to confirm architecture and composition including at least porosity, strut size, and pore size and to verify that printing or solvent processing does not introduce undesired compositional changes; conditioning the scaffold for cell interaction by preconditioning in serum containing media prior to cellular assays to verify biocompatibility and adhesion; and accepting the scaffold for biological use upon confirmation of architectural conformance and stable composition suitable for cell adhesion and tissue formation. In some embodiments, the method further comprises plasma treatment of the printed scaffold prior to cell seeding to increase wettability, wherein increased plasma exposure time reduces water contact angle and at five minutes enables complete droplet penetration through the scaffold. In still other embodiments, the method further comprises gas sterilization by ethylene oxide, wherein tensile and compressive mechanical properties of the cartilage domain scaffold remain unchanged pre and post sterilization as a release criterion. In yet other embodiments, the method further comprises adsorbing an adhesion protein to a bone region surface after plasma treatment by incubating the scaffold in fibronectin solution, drying, and proceeding to dynamic cell loading in a custom chamber.

[0087] In certain embodiments, the method further comprising non destructive micro computed tomography to quantify microarchitecture by layer, including side pores, fiber width, pore height, and pore width, and to validate infill density against mass / volume estimates across multiple samples as a post print fidelity assay. In some embodiments, printer maintenance is triggered by micro CT-detected microscopic defects and includes an extruder “cold pull” cleaning protocol that reduces layers with defects from about 15.4% to about 1.51%. IN particular embodiments, the method further comprises selecting and applying surface treatments to the articular region to reduce friction, including chondroitin sulfate conjugation via NaOH functionalization followed by EDC NHS, and application of a PTFE coating, each validated by contact angle reduction and / or friction coefficient reduction relative to control. In additional embodiments, the method further comprises printing complex non planar anatomical shapes with sacrificial, water dissolvable support material to avoid micro abrasions on articular surfaces duringDocket 88800730-000525support removal. In some embodiments, the method further comprises establishing process parameters for printing PLA / PBAT scaffolds at approximately 200-220°C while performing any subsequent hydrogel deposition at room temperature or about 40°C to maintain print fidelity and biological compatibility.

[0088] In certain embodiments, acceptance criteria for prepared scaffolds are aligned to finite element model-driven target properties, including compressive yield stress thresholds for cartilage and bone domains and tensile ultimate strength at the bone-cartilage interface. In some embodiments, the method further comprises in vitro degradation assessment demonstrating minimal mass loss and morphological stability over months at 37°C and in vivo integrity studies showing small mass loss (<5%) and no major morphology change with preserved compressive yield strength, thereby confirming suitability of the prepared scaffold for biological use. In other embodiments, prior to cell loading, the scaffold is incorporated into a cell seeding workflow using custom chambers for second or first components that enable dynamic loading of bone cells at defined densities under orbital agitation to promote uniform cell attachment. In yet other embodiments, the method further comprises integrating risk based design control activities, including identification of potential hazards and execution of nonclinical testing such as degradation, cytotoxicity, particulate formation, and implantation, with redesign triggered if results are unacceptable. In other embodiments, the method further comprises verifying layered implant architecture and material assignment in a multilayer scaffold (low friction coating, cartilage seeding layer, impermeable layer, bone seeding layer) as part of the preparation and release process for anatomical constructs. In additional embodiments, the method further comprises ex situ mechanical verification of prepared anatomical constructs, including compressive loading of the assembled joint to high loads with smooth, continuous forcedisplacement curves and absence of gross damage on pre / post scans prior to cadaveric testing.

[0089] The present disclosure additionally provides a quality-control method for compositional analysis of 3D printed polymer or polymer-ceramic filaments and detection of contamination, comprising: acquiring attenuated total reflectance Fourier transform infrared (ATR FTIR) spectra of raw materials, intermediate pellets, and printed filaments; preprocessing the spectra by baseline correction and normalization; extracting spectral features corresponding to polymer and ceramic constituents, including PLA and PBAT band signatures and phosphate peaks associated with 0 TCP and HA; comparing extracted features to a reference spectral library derivedDocket 88800730-000525from qualified inputs and prior acceptable prints to (i) quantify composition and verify no unanticipated changes due to solvent processing, melt blending, or printing and (ii) flag unexpected peaks or out of tolerance residuals as contamination; and outputting a pass / fail decision and recommended corrective actions for the filament production lot.

[0090] In addition, the present disclosure provides a method for introducing cell suspensions to inoculate anatomically shaped scaffold constructs for a biologic equivalent of a component of a conventional non-knee implant, the method comprising: (a) providing a three-dimensional anatomically shaped scaffold corresponding to at least one of a first component or a second component, the scaffold comprising a flexible, degradable polymer or polymer-ceramic composite with an internal porous architecture; (b) preparing a cell suspension comprising autologous adipose-derived stem cells (ASCs) and / or allogeneic human induced pluripotent stem cell (hiPSC)-derived chondrogenic or osteogenic cells; (c) mixing the cell suspension with a hydrogel carrier comprising HAMA or a genipin-crosslinked fibrin hydrogel (FibGen), optionally heparin-conjugated to promote lubricin adsorption and boundary lubrication; (d) loading the cellcontaining hydrogel into a bioprinter having a modified dual-channel nozzle and dispensing the mixture under controlled feed rate and nozzle height to deposit onto the scaffold surface; (e) inoculating the anatomically shaped porous construct by axial deposition and controlled infiltration of the cell-containing hydrogel to a predetermined depth within the pores; (f) inoculating the anatomically shaped scaffold by conformal deposition of the cell-containing hydrogel over an articular surface region and, optionally, into a bone-interfacing region, while maintaining a continuous coating thickness; (g) controlling coating thickness and infiltration depth by adjusting the dispensing feed rate and nozzle standoff height; and (h) verifying post-inoculation distribution and viability by imaging and mechanical adequacy of the coating, and releasing the construct for subsequent culture or implantation.

[0091] The present disclosure further provides a method for introducing cell suspensions to inoculate cylindrically shaped scaffold constructs for a biologic equivalent of a component of a conventional non-knee implant, the method comprising: (a) providing a three-dimensional cylindrically shaped scaffold corresponding to at least one of a first component or a second component, the scaffold comprising a flexible, degradable polymer or polymer-ceramic composite with an internal porous architecture; (b) preparing a cell suspension comprising autologous adipose-derived stem cells (ASCs) and / or allogeneic human induced pluripotent stem cell (hiPSC)-Docket 88800730-000525derived chondrogenic or osteogenic cells; (c) mixing the cell suspension with a hydrogel carrier comprising HAMA or a genipin-crosslinked fibrin hydrogel (FibGen), optionally heparin-conjugated to promote lubricin adsorption and boundary lubrication; (d) loading the cellcontaining hydrogel into a bioprinter having a modified dual-channel nozzle and dispensing the mixture under controlled feed rate and nozzle height to deposit onto the scaffold surface; (e) inoculating the cylindrically shaped porous construct by axial deposition and controlled infiltration of the cell-containing hydrogel to a predetermined depth within the pores; (f) inoculating the cylindrically shaped scaffold by conformal deposition of the cell-containing hydrogel over an articular surface region and, optionally, into a bone-interfacing region, while maintaining a continuous coating thickness; (g) controlling coating thickness and infiltration depth by adjusting the dispensing feed rate and nozzle standoff height; and (h) verifying post-inoculation distribution and viability by imaging and mechanical adequacy of the coating, and releasing the construct for subsequent culture or implantation.

[0092] Additionally, the present disclosure provides a method of introducing cell suspensions to inoculate three-dimensional printed scaffolds configured as cylindrical test coupons and anatomically shaped non-knee components for biological use as a biologic equivalent of a component of a conventional non-knee implant, the method comprising: providing a porous three-dimensional printed scaffold having a pore architecture and infill selected for cell seeding and load-bearing under non-knee-relevant loads; preparing a cell suspension comprising stem cell-derived bone cells for bone-domain regions and stem cell-derived cartilage cells for cartilagedomain regions; introducing the cell suspension into the scaffold to permeate interconnecting pores by at least one of direct perfusion or dispensing a cell-containing hydrogel slurry that is extruded to infiltrate the scaffold interstices; and incubating the inoculated scaffold under sterile culture conditions sufficient to promote cell adhesion and viability as confirmed by live / dead imaging. In certain embodiments, the method further comprises preconditioning and seeding bone-domain regions by plasma treating the scaffold to increase wettability, adsorbing fibronectin onto the bone region in a custom chamber, and loading a bone-cell suspension at 10 million cells per milliliter from a bottom port followed by orbital agitation at 120 rpm for three hours at 37°C in 5% CO2.In some embodiments, the anatomically shaped first and second components are seeded in dedicated chambers configured for each component to deliver the cell suspension into the intended bone-domain regions. In other embodiments, cartilage-domain inoculation comprises mixingDocket 88800730-000525cartilage cell clusters with methacrylated hyaluronic acid (HAMA, 1%) and LAP photoinitiator (0.03%) at 20 million cells per milliliter, submerging the cartilage region of the component for one minute, and UV-curing for two minutes at 405 nm to immobilize the cells in situ. In additional embodiments, seeding efficiency is increased by using higher cell-loading densities (5-10 million cells per milliliter) under gentle orbital agitation to minimize cell loss from the scaffold.

[0093] In yet other embodiments, cylindrical osteochondral constructs are inoculated and cultured up to 28 days with or without applied reciprocal shear loading to evaluate maintenance of mechanical properties and matrix elaboration during culture. In further embodiments, seeded anatomical implants are verified under ex situ compression with smooth, continuous forcedisplacement curves over repeated loads, indicating no failure of the inoculated constructs during culture. In some embodiments, the method further comprises validating print fidelity of seeded scaffolds by micro-computed tomography analysis of side pores, fiber width, pore height, and pore width, and corroborating infill density by mass / volume calculations across multiple printed samples as a post-print fidelity assay. In other embodiments, preparation for cell inoculation includes plasma treatment of the scaffold to reduce water contact angle and enable full droplet penetration through the porous architecture at five minutes of plasma exposure. In certain embodiments, the method further comprises design-control-aligned acceptance criteria for inoculated constructs, including target compressive yield stresses for cartilage and bone regions and minimum tensile strength at the cartilage-bone interface. In further embodiments, inoculated scaffolds are assessed for in vivo integrity demonstrating minimal mass loss (<5%) and no major morphology change with maintained compressive yield strength, confirming suitability of the inoculated construct for biological use. In yet further embodiments, cartilage-region inoculation is coordinated with multi -material printed layered scaffolds comprising a low-friction coating, a cartilage-seeding layer, an impermeable layer, and a bone-seeding layer, and cell seeding is directed to layer-specific regions. In still further embodiments, anatomical, non-planar constructs are printed with a water-dissolvable support to preserve articular surface smoothness for subsequent cell inoculation. In yet other embodiments, the inoculation workflow is documented in a manufacturing process overview from scaffold printing and cell generation through cartilage and bone cell seeding steps and final incubations prior to load / pack for transport. In particular embodiments, the method further comprises osteogenic differentiation of adipose-derived stemDocket 88800730-000525cells prior to bone-region seeding and chondrogenic differentiation for cartilage-region seeding, followed by chamber-based inoculation consistent with the defined process parameters.

[0094] The present disclosure also provides a device for seeding cells into porous polymer constructs, with a lower scaffold region and upper scaffold region separated by an impermeable layer, in a spatially-dependent manner, comprising: one or more inner walls, wherein the one or more inner walls form a container that conforms to the shape of the construct; a form fitting gasket / barrier separating the upper scaffold region and the lower scaffold region wherein the impermeable layer delineates the upper scaffold region and the lower scaffold region; one or more ports for introduction and evacuation of media with and without cells into the upper scaffold region and the lower scaffold region, and configured to enable introduction of media ± cells to the lower scaffold region and the upper scaffold region independently. In certain embodiments, cells are added to one of the upper scaffold region or the lower scaffold region. In other embodiments, cells, of a same or different type, are added to both the upper scaffold region or the lower scaffold region. In some embodiments, the method further comprising one or more perfusion ports, wherein the perfusion ports are configured to enable mixing or perfusion or to incorporate stir bars for media mixing.

[0095] Additionally, the present disclosure provides a device for seeding cells into a porous, three dimensional printed polymer construct that is a biologic equivalent of a component of a conventional non-knee implant, the device comprising: a fixture configured to receive and position a patient specific first or second scaffold having separate cartilage and bone domains; and a dispensing assembly configured to deliver hydrogels, cells, and biomaterials in a spatially dependent manner onto a first scaffold layer associated with bone and onto a second scaffold layer associated with cartilage. In some embodiments, the scaffold comprises a layered construct with an impermeable interface between a cartilage seeding region and a bone seeding region to prevent cell migration between domains. In other embodiments, the device further comprises a second seeding chamber having a bottom loading port, the chamber being configured to direct a bone cell suspension into the bone region while the implanted scaffold is held in position. IN additional embodiments, the device further comprises a first seeding chamber configured with a lid and fixture geometry to position the first component and deliver a bone cell suspension into the bone region. In other embodiments, the dispensing assembly is configured to deliver a cartilage cell hydrogel formulation to the cartilage region and includes a UV source to crosslink methacrylatedDocket 88800730-000525hyaluronic acid after submersion exposure. In further embodiments, the chamber is further configured to operate with plasma treated scaffolds and to perform fibronectin adsorption on the bone region prior to loading, thereby improving wettability and adhesion.

[0096] In particular embodiments, the chamber is configured for dynamic loading using an orbital shaker to maintain gentle agitation during seeding at cell loading densities between 5 and 10 million cells per milliliter. In certain embodiments, the dispensing assembly comprises a multi material printhead or nozzle system that assigns different materials to cartilage and bone regions during a single continuous operation on anatomical constructs. In some embodiments, the layered construct comprises, in order, a low friction rectilinear coating layer, a 70% infill alternating triangular cartilage seeding layer, a 100% infill impermeable layer, and a 70% infill alternating triangular bone seeding layer comprising polymer ceramic composite. In other embodiments, the device further comprises a quality control station configured for micro computed tomography analysis to quantify side pores, fiber width, pore height, and pore width by layer, and to validate infill density against mass / volume measurements as a post print fidelity assay. In yet other embodiments, quality control further includes monitoring and reduction of microscopic print defects through extruder “cold pull” cleaning documented by micro computed tomography before / after analysis. In still other embodiments, anatomical non planar components are printed and supported with a water dissolvable support to preserve articular surface smoothness prior to chamber based seeding. In additional embodiments, acceptance criteria for seeded constructs are aligned to finite element model derived targets for compressive yield stress for cartilage and bone regions and tensile ultimate strength at the cartilage-bone interface. IN particular embodiments, the device further comprises provisions for ex situ mechanical verification of seeded anatomical constructs under compressive loading with smooth, continuous force-displacement curves over repeated loads during culture.

[0097] In certain embodiments, the processor or control software documents the end to end manufacturing workflow including polymer printing, cell generation, cartilage and bone seeding steps in chambers, and incubation prior to transport. In some embodiments, the cartilage and bone regions and their fiber overlay patterns are predefined in CAD and used by the dispensing assembly to achieve region specific seeding in cylindrical test coupons and anatomical components. In further embodiments, the seeding operation for cartilage is performed by submerging the cartilage region in the cell laden HAMA solution for about one minute followed by UV curing at about 405Docket 88800730-000525nm for about two minutes. In other embodiments, bone cell seeding is performed at about 10 million cells per milliliter via the bottom loading port, followed by orbital agitation at about 120 rpm for about three hours at about 37°C and 5% CO2. In some embodiments, plasma treatment durations include increments up to about five minutes to enable full droplet penetration through the scaffold architecture before fibronectin adsorption. In yet other embodiments, the impermeable interface is verified within the layered construct during manufacturing and prior to seeding to prevent cell cross migration between cartilage and bone seeding regions.

[0098] The present disclosure also provides a scaffold for a biologic equivalent of a component of a conventional non-knee implant, comprising: an articular surface region configured to present a low-friction interface, the articular surface region including: (a) a hydrogel coating comprising HAMA or a genipin-crosslinked fibrin hydrogel (FibGen) conjugated with heparin to promote boundary lubrication by lubricin adsorption; and (b) a controlled infiltration of the hydrogel into a porous polymer scaffold underlying the articular surface region, wherein the scaffold further comprises: (c) a surface-smoothing treatment configured to reduce microabrasions on the articular surface region resulting from fabrication supports; and (d) a process control of hydrogel application comprising at least one of a defined nozzle height and a defined feed rate to achieve a target hydrogel layer thickness and infiltration depth, and wherein, when the scaffold is immersed in synovial-like fluid, the hydrogel coating provides interstitial fluid pressurization that bears a majority of an applied compressive load and lowers an effective friction coefficient at the articular surface region. In some embodiments, the scaffold further comprises controlling the directionality of the alignment of the 3D printed surface with the direction of sliding / movement. In other embodiments, the scaffold further comprises a polymer surface including conjugated glycosaminoglycans to act as a lubricant and reduce friction. In certain embodiments, the surface is coated with PTFE.

[0099] In some embodiments, the heparin conjugation density is selected to enhance lubricin adsorption relative to non-conjugated HAMA or FibGen, as evidenced by increased Zone 1 compressive modulus without loss at Zones 2-3. In other embodiments, the hydrogel layer thickness is controlled to a range correlated to feed rate. In yet other embodiments, a nozzle height of about 0.5-1.0 mm is used to achieve uniform coating over pores of about 300 pm. In still other embodiments, supports consist of PVA removable in water to prevent micro abrasions on theDocket 88800730-000525articular surface. In further embodiments, the hydrogel coated porous scaffold exhibits a dynamic compressive modulus within a target range for cartilage of about 15-60 MPa.

[0100] Additionally, the present disclosure provides a method of reducing friction of an articular region of a three-dimensional printed scaffold that is a biologic equivalent of a component of a conventional non-knee implant, comprising: preparing a fibrin-based hydrogel comprising genipin crosslinking and heparin conjugation; depositing the hydrogel as a surface coating on a scaffold articular region to promote boundary lubrication via lubricin adsorption and interstitial fluid pressurization; and controlling coating homogeneity and thickness by adjusting bioprinting parameters to produce a uniform, thin layer over the articular region. In certain embodiments, the surface treatment further comprises chondroitin sulfate conjugation of a PLA PBAT scaffold by NaOH functionalization, EDC NHS coupling, and overnight ChS submersion, which reduces water contact angle and friction coefficient relative to a control. In some embodiments, the articular region is subsequently coated with a PTFE spray that reduces friction coefficient relative to control and ChS conjugated samples. In additional embodiments, the layered scaffold further includes a 70% infill rectilinear coating layer selected to minimize friction at the articulating surface.

[0101] In further embodiments, hydrogel printing uses a modified dual printing unit and a two channel exit Y nozzle, and coating thickness and infiltration are controlled by stepper motor feed rate and nozzle height, yielding homogeneous thin hydrogel layers. In yet other embodiments, friction testing of the hydrogel coated scaffold demonstrates a friction coefficient that drops from about 0.095 to about 0.055 and remains stable over at least 100 cycles with minimal fluctuation. In particular embodiments, the hydrogel formulation is configured to support interstitial fluid pressurization and boundary lubrication via lubricin adsorption upon exposure to synovial fluids. In some embodiments, anatomical non planar articular shapes are printed with water dissolvable support to preserve surface smoothness prior to hydrogel coating. In certain embodiments, the articular coating step is incorporated into a manufacturing workflow that includes CAD definition of layered regions and anatomical surfaces to guide targeted friction reducing treatments on the cartilage region. In additional embodiments, the method further comprises micro CT assessment of post treatment surface region fidelity and pore level architecture to confirm uniform hydrogel coverage and absence of support removal damage. In further embodiments, the hydrogel coated, surface treated constructs maintain smooth, continuous force displacement curves withoutDocket 88800730-000525discontinuities during repeated loading in culture, indicating preservation of frictional performance without damage. In some embodiments, coating parameters and acceptance criteria are tied to finite element driven target properties and durability testing, including friction and wear evaluation in a loader bioreactor under physiologic shear loading.

[0102] The present disclosure further provides an articular-surface-treated non-knee scaffold comprising: a porous polymer or polymer blend lattice; and a hydrogel coating on an articular region, wherein the coating comprises genipin crosslinked fibrin formulated to promote lubricin adhesion for boundary lubrication and is deposited as a homogeneous thin layer tuned by feed rate and nozzle height.

[0103] The present disclosure also provides a system for producing a low-friction articular surface on a biologic non-knee implant component, comprising: a bioprinting unit configured to dispense a fibrin genipin hydrogel onto an articular region of a scaffold; and a controller configured to adjust feed rate and nozzle height to control coating thickness and homogeneity so as to reduce friction by enhancing interstitial fluid pressurization and lubricin-mediated boundary lubrication. In certain embodiments, the controller enforces acceptance criteria keyed to layer specific friction reducing architecture including an articular low friction coating layer over a cartilage seeding layer, an impermeable layer, and a bone seeding layer, thereby preserving low friction performance post assembly.

[0104] Additionally, the present disclosure provides a patient-specific implantation guide and fixation system for a layered biologic equivalent of a component of a conventional non-knee implant, comprising: (a) a custom surgical guide body defined by a computer-aided design (CAD) model generated from scan-derived geometry of the conventional non-knee implant component and corresponding patient bone interface regions; (b) one or more guide features that constrain bone resection to produce flat mating surfaces at the first and / or second interfaces consistent with the biologic implant’s scaffold design; (c) thickness control features that reference an articular layer and a bone scaffold layer of the biologic implant and define cut planes or stop surfaces so that the prepared host bone accommodates a target cartilage-layer thickness and a complementary bone-layer seating depth; (d) an output comprising digital tool-paths or positional data compatible with robotic or computer-assisted surgery for placement of the biologic implant in a single, continuous 3D-printed form without post-print assembly seams; and (e) a fixation technique configured for non-permanent fixation, comprising press-fit seating of the biologic implant’sDocket 88800730-000525porous scaffold against the prepared flat mating surfaces to promote osteointegration, wherein the custom surgical guide body, the thickness control features, and the robotic or computer-assisted placement collectively align the articular surface region to reduce peak contact pressure under physiologic loading, and the fixation technique avoids cemented or permanent hardware while enabling osteoconductive integration of the bone scaffold layer.

[0105] The present disclosure further provides a device for guiding implantation of a patient specific biologic non-knee arthroplasty component comprising: a patient specific positioning fixture configured to seat a layered first or second component that includes separate cartilage and bone scaffold domains; and guide surfaces configured to align the component for placement against surgically prepared flat mating bone interfaces using existing non-knee implantation methods and tools, optionally robotic, to achieve press fit contact at the bone interfaces. In certain embodiments, the device comprises a custom osteotomy cut guide configured to produce an additional first cut to accommodate a central slab that carries four layers throughout the implant. In some embodiments, the cut guide includes patient specific mating features and orientation references corresponding to top and bottom guide views to align the guide on the distal bone. In other embodiments, the device further comprising second and first jigs dimensioned for NOVAJoint (non-knee) components defined by in silico deformation and layered CAD renderings, to guide intraoperative seating and orientation of cartilage and bone domains. In additional embodiments, the device is configured for implantation using standard TKA robotic techniques as a standardized surgical method for placement and alignment.

[0106] In other embodiments, the device includes alignment features keyed to an implant architecture comprising a 70% infill rectilinear coating layer, a 70% infill alternating triangular cartilage seeding layer, a 100% infill impermeable layer, and a 70% infill alternating triangular bone seeding layer with hydroxyapatite and 0 tricalcium phosphate. In some embodiments, the device further comprises ancillary bioabsorbable fixation instrumentation compatible with 2.4 mm SmartNail poly 96L / 4D lactide copolymer nails and 2.0 mm TRIM IT enhanced PLLA pins. In certain embodiments, the fixation technique provides initial graft recipient compression comparable to headed screws and shear load to failure meeting levels of standard of care fixation devices for osteochondral shell procedures. In yet other embodiments, non planar anatomical constructs are printed with water dissolvable support material to preserve articular surface smoothness that interfaces with the guide. In still other embodiments, component positioningDocket 88800730-000525achieved using the device is validated ex situ under compressive loads up to ten body weights with smooth, continuous force-displacement curves and without gross damage prior to cadaveric testing. In additional embodiments, the device further comprises quality assurance features specifying target mechanical acceptance criteria aligned to finite element model derived thresholds, including compressive yield stress for cartilage and bone domains and minimum tensile strength at the interface. In particular embodiments, canine specific guides and fixation constructs are provided for cadaveric and in vivo canine evaluations with robotic testing to confirm anatomic positioning, alignment, and stability across cycles. In some embodiments, clinical use of the device is integrated into a standardized surgical method and post implant rehabilitation protocol in a phase 1 safety study design. In other embodiments, intraoperative verification aided by the device targets the absence of loosening, displacement, or nonunion and preservation of range of motion thresholds during cadaveric and clinical protocols. In further embodiments, the guide geometry is generated from patient specific CAD dimensions and sliced layered regions of the first and second components to ensure proper slab accommodation and layer alignment.

[0107] This disclosure is embodied in various forms illustrated in a set of accompanying illustrative drawings. Variations are contemplated as being a part of this disclosure, limited only by scope of various claims recited herein.BRIEF DESCRIPTION OF DRAWINGS

[0108] The set of accompanying illustrative drawings shows various exemplary embodiments of this disclosure. Such drawings are not to be construed as necessarily limiting this disclosure. Like numbers and / or similar numbering scheme can refer to like and / or similar elements throughout.

[0109] FIG. 1A, FIG. IB, FIG. 1C and FIG. ID. Rendering of the tibial component micro-CT scan using 3D slicer.

[0110] FIG. 2. CAD model of the knee components based on micro-CT scans.

[0111] FIG. 3 A and FIG. 3B. 3D printing of CAD models of knee components.

[0112] FIG. 4A and FIG. 4B. Maximum loads in implant tibial components before (FIG.4A) and after (FIG. 4B) in silico deformation.

[0113] FIG. 5. Overview of full manufacturing workflow.

[0114] FIG. 6. Overview of drug product manufacturing.Docket 88800730-000525

[0115] FIG. 7. Porous biomaterials printed using either gyroid infill or alternating infill pattern (90, 45, 135, 0 degrees).

[0116] FIG. 8 A, FIG. 8B, FIG. 8C, FIG. 8D, FIG. 8E, FIG. 8F, FIG. 8G, FIG. 8H, FIG. 81, FIG. 8J, FIG. 8K, FIG. 8L, FIG. 8M, FIG. 8N, FIG. 80, FIG. 8P, FIG. 8Q, FIG. 8R and FIG. 8S. CAD models of various porous structures. Alternating pattern with various infills (FIG. 8A, FIG.8B, FIG. 8C, FIG. 8D). Honeycomb pattern with various infills (FIG. 8E, FIG. 8F, FIG. 8G, FIG, 8H). Rectilinear pattern (FIG. 81), triangle pattern (FIG. 8J), gyroid pattern (FIG. 8K), Hilbett curves pattern with 25%, 50% and 70% infill, (FIG. 8L, FIG. 8M, FIG. 8N, FIG. 80), Archimedean cords pattern with 25%, 50% and 70% infill (FIG. 8P, FIG. 8Q, FIG. 8R, FIG. 8S).

[0117] FIG. 9A, FIG. 9B, FIG. 9C and FIG. 9D. CAD models of alternating infill pattern.100% infill (FIG. 9A), 70% infill (FIG. 9B), 50% infill (FIG. 9C) and 25% infill (FIG. 9D).

[0118] FIG. 10A and FIG. 10B. CAD models of various porous structures. Porous triangular infill (FIG. 10A) and porous rectilinear infill (FIG. 10B).

[0119] FIG. 11A and FIG. 11B. Studies on CAD models of various stacked patterns. Alternating layers using the triangular pore shape (FIG. 11 A and FIG. 1 IB).

[0120] FIG. 12A, FIG. 12B, FIG. 12C and FIG. 12D. Implant geometry design. Modularity was removed from a CAD model of a knee implant and smooth surfaces were created (FIG. 12A). Further modification involved introducing separate articular and bone scaffold domains for both the femoral and tibial components of the implant (FIG. 12B). FEM of Model 1 implant. The von Mises (effective stress of the Model 1 implant is shown in FIG. 12C, and the contact pressure of the Model 1 implant is shown in FIG. 12D.

[0121] FIG. 13 A and FIG. 13B. FEM on initial scans of commercial knee. FEM setup for the applied force of 10 body weights is shown in FIG. 13A. The results are shown in FIG. 13B.

[0122] FIG. 14AandFIG. 14B. In silico plastic deformation performed in the commercial knee scan models. FEM results shown in FIG. 14A and FIG. 14B.

[0123] FIG. 15A and FIG. 15B. FEM of one embodiment of presently disclosed biologic living TKA joint. Contact pressure before altering tibial geometry (FIG. 15A) and after altering tibial geometry (FIG. 15B).

[0124] FIG. 16A, FIG. 16B and FIG. 16C. Model of cartilage domain layer on the femoral component.

[0125] FIG. 17A and FIG. 17B. Model of cartilage domain layer on the tibial component.Docket 88800730-000525

[0126] FIG. 18A and FIG. 18B. Failure testing of one embodiment of the presently disclosed biologic living TKA joint. Pictures of the tibial loading jig, the femur loading jig, the tibial component and the femoral component before the test is shown in FIG. 18A, and after the test is shown in FIG. 18B.

[0127] FIG. 19. Picture of implant and respective jig components configured for mechanical testing.

[0128] FIG. 20. 3D printing setup for NOVAKnee scaffold with four distinct regions. A sacrificial support structure is also included for printing.

[0129] FIG. 21. 3D printing setup for NOVAKnee scaffold with three distinct regions, including fiber overlay pattern for each implant region.

[0130] FIG. 22A and FIG. 22B. A representative NOVAKnee femoral scaffold with a slab (FIG. 22A) and an example of a femur cut for a traditional knee implant with the approximate location of the additional cut necessary for NOVAKnee indicated in green (FIG. 22B).

[0131] FIG. 23. CAD renderings of scanned traditional TKA implant components.

[0132] FIG. 24A and FIG. 24B. Different depictions of one embodiment of femoral component of NOVAKnee. FIG. 24A. CAD renderings of scanned NOVAKnee scaffold designed through in silico deformation. FIG. 24B. CAD renderings of femoral component of NOVAKnee scaffolds with exploded views of the sliced regions of the femoral component.

[0133] FIG. 25. CAD renderings of scanned traditional TKA implant components.

[0134] FIG. 26A and FIG. 26B. Different depictions of one embodiment of tibial component of NOVAKnee. FIG. 26A. CAD renderings of scanned traditional TKA implant components. FIG. 26B. CAD renderings of tibial components of NOVAKnee scaffolds with exploded views of the sliced regions of the tibial component.

[0135] FIG. 27A, FIG. 27B and FIG. 27C. One embodiment of 3D printed multilayer NOVAKnee construct. FIG. 27A. Full construct. FIG. 27B. Femoral component. FIG. 27C. Tibial component.

[0136] FIG. 28 A and FIG. 28B. CAD rendering of layered femoral component. FIG. 28 A. Side view of layered femoral component. FIG. 28B. Side view of expanded portion of layered femoral component.Docket 88800730-000525

[0137] FIG. 29A and FIG. 29B. FEM analysis maximum loads for commercial knee implant. FIG. 29A. Contact pressure (compression). FIG. 29B. Maximum principal stress (tension).

[0138] FIG. 30A and FIG. 30B. FEM analysis maximum force measurements for in silico deformation designed NOVAKnee. FIG. 30A. Contact pressure (compression). FIG. 30B. Maximum principal stress (tension).

[0139] FIG. 31 A, FIG. 3 IB, FIG. 31C, FIG. 3 ID and FIG. 3 IE. Representative overall dimensions of NOVAKnee sized to be similar to a size 9E of a commercially available conventional knee implants. FIG. 31 A. Side view of femoral component. FIG. 3 IB. Top view of femoral component. FIG. 31C. Bottom view of femoral component. FIG. 3 ID. Side view of femoral component. FIG. 3 IE. Front view of femoral component.

[0140] FIG. 32A, FIG. 32B, FIG. 32C and FIG. 32D. Representative overall dimensions of NOVAKnee sized to be similar to a size 9E of a commercially available conventional knee implants. FIG. 31 A. top view of tibial component. FIG. 3 IB. Bottom view of tibial component. FIG. 31C. Front view of tibial component. FIG. 3 ID. Side view of tibial component.

[0141] FIG. 33A, FIG. 33B and FIG. 33C. One iteration of cut guide for one embodiment of NOVAJoint design. FIG. 33 A. Image of femur showing location of additional cut for central slab. FIG. 33B. Top view of cut guide. FIG. 33B. Bottom view of cut guide.

[0142] FIG. 34. 3D Printed multilayer osteochondral scaffolds. Example of a two-layer tibial implant component.

[0143] FIG. 35A, FIG. 35B, FIG. 35C and FIG. 35D. Anatomical knee 3D printing. Complex non-planar shapes (FIG. 35 A) require the use of support materials during 3D printing (yellow portion of FIG. 35 A). Physical removal of the support surfaces led to micro-abrasions and a roughened surface (FIG. 35B). To circumvent this, a water dissolvable material support material was used (FIG. 35C) and print orientation on build plate was optimized. The resultant process for 3D printed multilayer osteochondral scaffolds for the tibial and femoral component is shown in FIG. 35D.

[0144] FIG. 36A, FIG. 36B, FIG. 36C and FIG. 36D. Representative pCT of 3D printed porous scaffold microarchitecture (triangular infill). FIG. 36A. Side pores. FIG. 36B. Fiber width. FIG. 36C. Pore height. FIG. 36D. Pore width.

[0145] FIG. 37. Modified 3D printed Y-shaped nozzles.Docket 88800730-000525

[0146] FIG. 38A, FIG. 38B, FIG. 38C and FIG. 38D. Modified dual printing unit. Prototype (FIG. 38A) and final version (FIG. 38B, FIG. 38C and FIG. 38D).

[0147] FIG. 39A and FIG. 39B. Different types of Y-nozzles. FIG. 39A - screw-type. FIG. 39B - two-channel exit.

[0148] FIG. 40. Process used to deposit cartilage and hydrogel solution onto cartilage regions of NOVAKnee-T.

[0149] FIG. 41A, FIG. 41B, FIG. 41C and FIG. 41D. 3D printed bone domain scaffolds. Porous grid structure on first scale (FIG. 41 A), and magnification of porous grid structure on first scale (FIG. 4 IB). Porous grid structure on second scale (FIG. 41C), and magnification of porous grid structure on second scale (FIG. 4 ID).

[0150] FIG. 42A and FIG. 42B. Additional 3D printed bone composite scaffolds. PCL-10% ZnO (FIG. 42 A), and PCL-10%ZnO+20% 20 / 80 HA / TCP (FIG. 42B).

[0151] FIG. 43. Optical images and SEM images of PCL (top panel) and PCL-ceramic (bottom panel) specimens manufactured with the bioprinter.

[0152] FIG. 44A, FIG. 44B and FIG. 44C. Chamber used to seed bone cells onto tibial component. Rendering (FIG. 44A - exploded view; FIG. 44B - cutaway view) and photo (FIG.43C) of.

[0153] FIG. 45 A,. FIG. 45B, FIG. 45C and FIG. 45D. Chamber used to seed bone cells onto femoral component. Rendering of chamber lid (FIG. 45 A), side view (FIG. 45B) and femur loading chamber (FIG. 45C) and photo (FIG. 45D).

[0154] FIG. 46A and FIG. 46B. Renderings of NOVAJoint. FIG. 46A - tibial component. FIG. 46B - femoral component.

[0155] FIG. 47. 3D printing setup for NOVAKnee scaffold with two distinct regions, including fiber overlay pattern for each.

[0156] FIG. 48. CAD model of 3D printed scaffold structures.

[0157] FIG. 49. CAD design of cartilage (blue) and bone (pink) overlaid structure in a cylindrical test coupon, with varying layer thicknesses.

[0158] FIG. 50A, FIG. 50B, FIG. 50C and FIG. 50D. Implantation of NOVAJoint into cadaveric knee and mechanical testing.Docket 88800730-000525

[0159] FIG. 51A, FIG. 51B, FIG. 51C and FIG. 51D. Photographs of a NOVAKnee scaffold implanted in a cadaveric human knee (FIG. 51 A and FIG. 5 IB) and being tested biomechanically after implantation (FIG. 51C and FIG. 5 ID).

[0160] FIG. 52A and FIG. 52B. Renderings of canine NOVAKnee (FIG. 52A) and custom femoral osteotomy guide (FIG. 52B).

[0161] FIG. 53A, FIG. 53B, FIG. 53C and FIG. 53D. Implantation of canine NOVAKnee-T into canine joint. Guide placement (FIG. 53 A), femoral cut (FIG. 53B), tibial cut (FIG. 53C) and implantation (FIG. 53D).

[0162] FIG. 54A, FIG. 54B and FIG. 54C. Multilayer porous ball-and-socket joint scaffold in 3D printer slicing software. The non-porous impermeable layers in each component are in the same material as the cartilage domain. Shown are both components (FIG. 54A), the acetabular component - plastic liner(FIG. 54B) and the femoral head - hip stem (FIG. 54C).

[0163] FIG. 55A, FIG. 55B and FIG. 55C. 3D printed multilayer porous ball-and-socket joint scaffold. Blue material is bone domain. Pink material is cartilage domain (and impermeable layer). FIG. 55A - both components separate. FIG. 55B - both components mated. FIG. 55C. Both components separate, showing inner portion of socket.

[0164] FIG. 56. Close-up photographs of 3D printed scaffold porous structure. Blue material is bone domain. Pink material is cartilage domain (and impermeable layer).

[0165] FIG. 57. Images (left panel) and range of dimensions (right panel) available for proDISC spinal implant.

[0166] FIG. 58. 3D printed NOVASpine scaffolds in two different sizes and shapes.

[0167] FIG. 59. Image of proDISC spinal implant.

[0168] FIG. 60. Images of proDISC spinal implant showing range of motion.

[0169] FIG. 61. CAD models of rectangular and disc-like designs for NOVASpine implants.

[0170] FIG. 62A, FIG. 62B and FIG. 62C. Multilayer porous spinal disc implant scaffold in 3D printer slicing software. The non-porous impermeable layers in each component are in the same material as the cartilage domain. FIG. 62A - vertebral side. FIG. 62B - impermeable layer. FIG. 62C - cartilage side.

[0171] FIG. 63 A, FIG. 63B and FIG. 63C. 3D printed multilayer porous NOVASpine scaffolds made of various bioresorbable materials and printed with various infill densities. FIG.Docket 88800730-00052563 A - side view. FIG. 63B - cartilage. FIG. 63C - bone (top panel) and complete scaffold (bottom panel).

[0172] FIG. 64. Close-up photographs of 3D printed scaffold porous structure. Blue material is bone domain. White material is cartilage domain (and impermeable layer).DETAILED DESCRIPTION

[0173] The present disclosure is now described more fully with reference to the set of accompanying illustrative drawings, in which exemplary embodiments of this disclosure are shown. This disclosure can be embodied in many different forms and should not be construed as necessarily being limited to the exemplary embodiments disclosed herein. Rather, the exemplary embodiments are provided so that this disclosure is thorough and complete, and fully conveys various concepts of this disclosure to those skilled in a relevant art.

[0174] The present disclosure provides compositions and methods for preparation of a novel patient-specific entirely biologic living non-knee joint joint replacement. This biologic nonknee joint (for example hip, shoulder and spine) represents a permanent and final procedure for a ravaged arthritic joint, allowing normal function in <4-6 weeks postop. The presently disclosed biologic living non-knee joint is manufactured on-demand for the patient using robust fabrication methods (allogenic implant <24 hours and autologous implant <30 days), representing a confluence of the inventors cutting edge innovations in stem cell biology, biomaterials and tissue engineering technologies, and biomechanics. The presently disclosed biologic living non-knee joint replicates the strategy used in implant surgery that has press-fitted first and second components designed to perform together. Unlike conventional joints, the presently disclosed biologic living non-knee joint is entirely resorbable ( .g., no permanent fixation), non-immunogenic, chondroinductive, osteoinductive, and will recapitulate native tissue properties. The present disclosure provides an allogenic version suitable for the vast majority of patients, as well as an autologous version tailored for individuals who are immunocompromised or with more complicated medical histories. The hierarchical structure of the presently disclosed biologic living non-knee joint includes a 3D printed scaffold mesh of biocompatible / bioresorbable polymer having mechanical properties sufficient to withstand loads typical for human non-knee joint infused with a cell-seeded biocompatible hydrogel doped with spatially defined osteoinductive or chondroinductive factors that promote de novo bone and cartilage formation inDocket 88800730-000525situ post implantation. Together, the combination scaffold mesh-hydrogel recapitulates the natural load bearing and tribological properties of articular cartilage while fostering growth of new cartilage to take over this function in time.

[0175] The following examples are included to demonstrate illustrative embodiments of the present disclosure. It should be appreciated by those of skill in the art that the techniques disclosed in the examples that follow represent techniques discovered by the inventors to function well in the practice of the present disclosure, and thus can be considered to constitute one embodiment of modes for its practice. However, those of skill in the art should, in light of the present disclosure, appreciate that many changes can be made in the specific embodiments that are disclosed and still obtain a like or similar result without departing from the spirit and scope of the present disclosure.

[0176] EXAMPLE 1 - Development of Biologic Living TKA Joint

[0177] This example details the development of a biologic living TKA joint, that can be used in lieu of standard metal and ultrahigh-molecular weight polyethylene (UHMWPE) artificial knees, while also taking advantage of existing surgical implantation methods and tools, such as robotic surgery. The inventors opted to adapt existing artificial knee designs in the design of the presently disclosed biologic living TKA joint. This was done by modifying the geometry of the articular surfaces of the presently disclosed biologic living TKA joint design and selecting suitable polymer blends or composites for the porous scaffold that exhibit the right amount of porosity for cell -seeding, and the right amount of strength and toughness to sustain stresses under loading.

[0178] Knee replacements typically employ metal and ultrahigh-molecular weight polyethylene (UHMWPE). Metal components have the requisite strength to withstand knee loads that reflect activities of daily living, and decades-old clinical experience suggests that UHMWPE is also adequately suited for withstanding these loads. The geometry of a representative metal— plastic artificial knee joint was scanned, and the scanned models were modified by smoothing the surfaces interfacing with the patient bone. Then finite element analyses of these designs was performed to examine the level of stresses that would be achieved under activities of daily living, using “running” as the worst-case loading scenario (running produces peak knee loads that approach ten times body weight). In these finite element analyses, loading was first performed at full extension (0° flexion), from zero to 10 body weights (BW), which is reported in the literature as the peak loading in the knee during running. Then flexion up to 100° was performed with theDocket 88800730-000525compressive load decreasing to 7 BW in conformation with biomechanics analyses drawn from the literature. An internal-external rotation of ±3° was imposed to ensure that the NOVAKnee-T had enough laxity to prevent undue biomechanical outcomes upon implantation. The body weight was calculated based on the BMI of the population of men and women who undergo total knee implants. A safety factor of at least lx was built in for potential failure during compression, and greater than 2x for tension at 10 BW (forces significantly higher than will be experienced by patients complying with post-surgical protocols for activity restrictions). Loading up to 10 x for designing and testing components for human use; alternative weighting factor for testing, i.e., range from 3-15 x; loading at different factor for dog or other animal.

[0179] Through this finite element method (FEM) analysis, it was discovered that running loads on existing knee implant designs can produce stresses that exceed the yield stress of UHMWPE, which would cause permanent deformation of UHMWPE components. Observations from surgical retrievals of knee components during TKA revisions also suggest these permanent deformations in the UHMWPE can occur. Thus, it was concluded that UHMWPE must have remarkably resilient properties that allow it to deform permanently without causing it to fail (e.g., crack). In contrast, the presently developed 3D-printed bioresorbable polymer materials that meet the design goals for the presently disclosed biologic living TKA joint do not have a clinically demonstrated ability to deform permanently while maintaining sufficient toughness and resistance to fracture. Therefore, the presently disclosed biologic living TKA joint design must produce lower stresses than existing artificial knee joint designs in order for the 3D-printed scaffold to remain intact (uncracked) during activities of daily living.

[0180] Starting from modified scanned models (smoothed surfaces interfacing with the patient bone) of existing artificial knee joint designs, the inventors increased the congruence of the articular surfaces of the femoral component and tibial insert using finite element analyses with plastic deformation of the tibial insert under loading; that is, inducing permanent deformation in silica to develop the presently disclosed biologic living TKA joint designs. First, permanent deformations were induced while the knee is near full-extension and loaded under five body weights, then additional permanent deformations were induced by allowing some amount of internal / external rotation between the femoral component and the tibial insert to adequately replicate the native knee’s range of motion. Then these procedures were repeated by allowing permanent deformation of the femoral component instead of the tibial component to furtherDocket 88800730-000525enhance joint congruence under various loading conditions. All investigations involve different amounts of loads and internal / external moments at various flexion angles.

[0181] The micro-computed tomography (micro-CT) scans of metallic femoral and tibial components and UHMWPE tibial inserts introduce a considerable amount of imaging artifacts, requiring the use of customized segmentation and smoothing algorithms to construct computer models from those scans. The micro-CT scans of metallic tibial components were examined. A 3D model constructed from the micro-CT scans was prepared. The metallic femoral component and UHMWPE tibial insert were segmented into surface representation in Meshmixer. The rendering of the tibial component micro-CT scan using 3D slicer is shown in FIG. 1A, FIG. IB, FIG. 1C and FIG. ID. The reduction of metal artifacts from micro-CT scanning of the femoral component was resolved. The tibial insert is made of UHMWPE, and does not produce metal artifacts.

[0182] The rendered surfaces obtained from the micro-CT scans were imported into the SolidWorks CAD software, from which CAD models of the knee components were made (FIG. 2), which were then 3D printed (FIG. 3A and FIG. 3B).

[0183] The scans are not perfect replicas of the original shapes. These computer models are altered to produce flat-mating surfaces at the interface of the femoral and tibial component and patient bone, to promote better osteointegration of the presently disclosed biologic living TKA joint porous scaffold with the patient bones, whereas existing artificial knee designs may employ suitable recesses for accommodating porous metals that enhance osteointegration. The pore structure of the presently disclosed biologic living TKA joint 3D-printed components differs considerably from the smooth structure of the metallic femoral and tibial components of existing artificial knees. The presently disclosed biologic living TKA joint design is biodegradable, and its shape can thus evolve over time in vivo. At the completion of in silico plastic deformation of the presently disclosed biologic living TKA joint femoral and tibial surface, the shapes of the articular surfaces of the presently disclosed biologic living TKA knee joint differ even further from those of the original scans of existing artificial knee joints. The modified scanned models of implant femoral components are shown before (FIG. 4A) and after (FIG. 4B) the in silico deformation, which results in the presently disclosed biologic living TKA joint design having more congruent and lower maximum contact pressure.Docket 88800730-000525

[0184] An overview of the NOVAKnee manufacturing process is detailed below. In step 1A, biodegradable polymer scaffolds are 3D printed, in step IB, cartilage and bone cells are generated from adipose-derived stem cells (ASCs) or human induced pluripotent stem cells (iPSCs), in step 2, the ASC or iPSC cells are seeded onto the scaffold, and the final NOVAKnee product is generated. An overview of the full manufacturing workflow is shown in FIG. 5, and an overview of drug product manufacturing is shown in FIG. 6. The final drug product process details are shown below in Table 1.Table 1

[0185] EXAMPLE 2 - Implant Design

[0186] The present design strategy is inspired by existing TKA implants and associated robotic surgery for producing the presently disclosed biologic living TKA joint, taking advantage of current standard of care for TKA surgeries and capitalizing on the extensive track -record of success of existing TKAs, the design of which compensates for the removal of menisci in the treated knee joint. This strategy allows focusing on producing a bioresorbable-polymer-basedDocket 88800730-000525implant shape that can withstand human knee loads (up to 10-times body weight) without undergoing significant damage (such as large permanent deformations or cracking).

[0187] CAD models of porous structures were created, specifically models for specific porosities and pore structure. Studies were performed to determine the contributions of pore structure / shape on compressive properties of PLA by comparing gyroid to alternating pore structure. Porous biomaterials were printed using either gyroid infill or a custom designed alternating infill pattern (90, 45, 135, 0 degrees) using a Prusa MK4 with no perimeter layer (FIG. 7). Gyroid pattern has similar performance to alternating infill pattern. Porous cylinders (N=6 per group) were tested under equilibrium unconfined compression (2% strain applied at 0.0217 mm / s, based on D695 ASTM standard, after a 5N preload) which resulted in comparable moduli for both patterns (p=0.24 by t-test). Dynamic compressive testing was also performed with no significant effect of infill pattem / pore shape observed. The results indicate that pore shape is not a significant variable contributing to the compressive modulus of PLA. These findings obviate the need to model pore shape.

[0188] CAD models were prepared of various porous structures. Scaffold pore shape (alternating pattern, honeycomb) was produced using the Prusa Slicer software. The alternating pattern with various infills is shown in FIG. 8A, FIG. 8B, FIG. 8C and FIG. 8D, and the honeycomb pattern with various infills is shown in FIG. 8E, FIG. 8F, FIG. 8G and FIG. 8H. Other patterns (rectilinear (FIG. 81), triangle (FIG. 8J) and gyroid (FIG. 8K) were also prepared. Other patterns (Hilbett curves pattern with 25%, 50% and 70% infill, FIG. 8L, FIG. 8M, FIG. 8N and FIG. 80; Archimedean cords pattern with 25%, 50% and 70% infill, FIG. 8P, FIG, 8Q, FIG. 8R and FIG. 8S) were ruled out. Compressive dynamic modulus of the alternating pattern and the honeycomb pattern were determined to be comparable.

[0189] Multiple CAD representations of porous scaffold structures were generated using the 3D-Prusa printer software. A single-direction infill pattern was created, with infill directions 0°, 45° and 90° (to loading direction), 100% infill, and strut size 0.4 mm wide, 0.2 mm thick. An alternating infill pattern was created (0, 45, 90, 135, consecutively, and strut size 0.4 mm wide and 0.2 mm thick. 100% infill (FIG. 9A), 70% infill (FIG. 9B), 50% infill (FIG. 9C) and 25% infill (FIG. 9D).

[0190] CAD models were made of triangular (FIG. 10A) and rectilinear (FIG. 10B) porous structures. Triangular infill pattern creates triangular pores that are stacked on top of one anotherDocket 88800730-000525through the thickness of the scaffold. Rectilinear consists of layers that have layers alternating by 90°. The compressive equilibrium modulus of triangular pore shape was significantly better than the alternating infill pattern (90, 45, 135, 0 degrees) or rectilinear. Dynamic compressive testing showed the greatest modulus for triangular pattern. The triangular infill pattern improved compressive properties but poses a challenge for cellular communication.

[0191] In certain embodiments, the pores can be of a single shape or varying shapes, such as, for example, triangular, square, rectangular, pentagonal, hexagonal, heptagon, octagonal, nonagonal, decagonal, circular, oval, or any combinations of the foregoing. In certain embodiments, the pores may overlap from layer to layer. In certain embodiments, the pores interconnect to create a plurality of continuous passages through the device. In certain embodiments, a majority of the pores are interconnected.

[0192] Compressive properties of various pore geometries (PLA, 70% infill) were determined. The geometries studied were honeycomb infill pattern, triangular infill pattern (aligned triangles), triangular infill pattern (074574570°), triangular infill pattern (0° / 45o / 135o / 90°), rectilinear infill pattern (0° / 90°) and alternating infill pattern (O74579O71350). The results for yield stress, elastic modulus and toughness show that the aligned triangular pattern produces significantly higher yield stress (66.9±1.1 MPa) than the other patterns. However, this triangular pattern does not allow pore communication sideways.

[0193] Alternating layers (90745745790°) using the triangular pore shape (FIG. 11 A and FIG. 11B) and stacked layers using the triangular pore shape were analyzed. The results for compressive equilibrium modulus and dynamic modulus show the alternating layers and stacked layers result in comparable modulus, but the alternating pattern is more effective for cellular communication.

[0194] Various pore structures were tested. Thinner struts and thicker struts were prepared and tested. For the thinner struts, the compressive test / PLA print conditions were fast strain rate 80% / s, porosity 70% infill, alternating pattern, and nozzle size 0.4 mm. The extrusion parameters: width (XY) 0.4 mm, and thickness (Z) 0.2 mm. The mechanical properties: yield strength 45 + 1.7 MPa, elastic modulus 790 + 43 MPa, toughness 59.1 + 3.5 MPa. For the thicker struts, the compressive test / PLA print conditions were fast strain rate 80% / s, porosity 70% infill, alternating pattern, and nozzle size 0.4 mm. The extrusion parameters: width (XY) 0.8 mm, and thickness (Z) 0.6 mm. The mechanical properties: yield strength 48 + 0.7 MPa, elastic modulus 1040 + 13Docket 88800730-000525MPa, toughness 57.7 + 0.3 MPa. Increasing strut cross-sectional area buy 6X increased elastic modulus by 32% and yield strength by 7%.

[0195] Compressive testing of solid and porous polycaprolactone (PCL) polymers was conducted. The first sample was 100% infill, low strain rate (0.1% / s), the second sample was 70% infill, low strain rate (0.1% / s), and the third sample was 70% infill, high strain rate (80% / s). Porous PCL (70% infill) exhibits similar yield strength to solid PCL (100% infill) when loaded at nearphysiologic strain rates (physiologic strain rates are ~300% / s) without increasing Young’s modulus excessively.

[0196] PLA is viscoelastic, and therefore its properties may depend on the loading rate. Therefore, the tensile properties on solid PLA (100% infill, alternating rasterization (0°, 45°, 90°, 135°)) as a function of the elongation rate was tested, and an increase in failure stress as the elongation rate was increased from 0.06% to 0.6% / second was observed. Above 5 mm / second the failure stress plateaued, suggesting no further enhancement in strength with higher strain rates. The strain rate in vivo is estimated to be on the order of 300% / second during running. Fracture strength of solid polylactic acid (PLA) as a function of tensile strain rate (0.06 % / s, 0.1 % / s, 1% / s, 6% / s, 10% / s or 0.05 mm / s, 0.085 mm / s, 0.85 mm / s, 5 mm / s and 8.5 mm / s, respectively) shows strength increases statistically (p<0.01) from 42 MPa at 0.06% / s to 59 MPa at 6% / s. The strength plateaus with further increase of strain rate to 10% / s. Young’s Modulus of solid PLA as a function of strain rate 0.06 % / s, 0.1 % / s, 1% / s, 6% / s, 10% / s or 0.05 mm / s, 0.085 mm / s, 0.85 mm / s, 5 mm / s and 8.5 mm / s, respectively) does not show a consistent trend, suggesting that damage accumulation occurs at higher strain rates, prior to complete failure.

[0197] Since cells are seeded in the presently disclosed biologic living TKA joint scaffolds, the failure properties of porous polymers were studied. Strength decreases significantly (p<0.001) with increased porosity. Fracture strength of PLA as a function of sample porosity [0%, 30%, 50% and 75% (or 100%, 70%, 50%, and 25% infill, respectively] were tested at a strain rate of 0.06 % / s. Solid PLA meets the SF=3 tensile strength bone domain requirements determined by FEM of > 27 MPa, while all tested porosities fall below this value (though still near / above the 9 MPa peak stresses in the implant FEM for 30% and 50% porosity). Representative samples when tested at 1% / s exhibit brittle failure.

[0198] Tensile testing was performed on solid (100% infill) polymer blends to obtain the necessary mechanical properties for the presently disclosed biologic living TKA joint. DifferentDocket 88800730-000525blends of PLA and PBAT (Ecovio® T230B (>75% PLA + <25% PBAT), Ecovio® F2224 (45% PLA + 55% PBAT) and Ecovio® F2341 (10% PLA + 90% PBAT)) were tested, with PLA, PCL and PBAT used as controls. Elongation rate for PLA and Ecovio-T2308 was 0.05 mm / s, all others at 0.5 mm / s. Yield strength of polylactic acid (PLA), polycaprolactone (PCL), polybutylene adipate terephthalate (PBAT) and Ecovio® (PLA+ PBAT) blends showed that compared to 100% PLA, Ecovio® blends have decreased yield strength. Compared to 100% PBAT, Ecovio® blends have increased yield strength. Young’s modulus of PLA, PCL, PBAT, and Ecovio® (PLA + PBAT) blends showed that compared to 100% PLA, Ecovio® blends have decreased stiffness. Compared to 100% PBAT, Ecovio® blends have increased stiffness. Toughness of PLA, PCL, PBAT, and Ecovio® (PLA + PBAT) blends showed that compared to 100% PLA, Ecovio® blends have increased toughness. Compared to 100% PBAT, Ecovio® blends have decreased toughness. These solid PLA, PCL, PBAT, and Ecovio® blends all meets the SF=3 tensile strength bone domain requirements determined by FEM of > 6 MPa.

[0199] Crystallinity of 3D printed PLA / PBAT blends was determined. The blends tested were 90 / 10 PLA / PBAT, 45 / 55 PLA / PBAT, and 10 / 90 PLA / PBAT. The results are shown below in Table 2.Table 2

[0200] PLA / PBAT blends after 3D printing do not have a significant difference in crystallinity %. 3D printed scaffolds across various blends exhibit similar crystallinity and thermal response (Tg and Tm).

[0201] Compressive testing of solid PLA polymers was performed. Solid PLA specimens (100% infill, 12.5 mm diameter, 12.5 mm height) before and after compressive testing were examined. Stress-strain response of PLA in compression, tested on Instron materials testing system, showed yield stress is 68.7 + 0.8 MPa, ultimate stress (strength) is 102.9 + 1.5 MPa,Docket 88800730-000525Young’s modulus is 1.66 + 0.04 GPa, and toughness is 33.74^0.4 MPa. Contrary to its tensile response, PLA exhibits significant toughness in compression (ductile failure at high strains). PLA scaffolds are much more resilient to loading in compression than tension. These properties provide sufficient strength to withstand physiological loads, such as running.

[0202] Due to the large difference in melting temperatures of PLA and PCL, it is difficult to melt blend these two polymers together. However, a blend of 70% PLA and 30% PCL was tested for their compressive properties and compared to those of PCL and PLA. Results show that PLCL has favorable modulus and toughness but lower yield strength then PCL (for PLCL, yield strength = 6.5 ±1.0 MPa, Young’s modulus = 170 ± 14 MPa, and toughness = 20.1 ± 1.6 MPa). Nevertheless, PLCL exhibited a 93% recovery in thickness 24 hours after testing, whereas PCL exhibited permanent deformation.

[0203] Mechanical testing of polymer blends was performed. Compressive properties of PBAT and PLA + PBAT blends were studied. The sample / compressive test parameters for PBAT was 100% infill, high strain rate (80% / s), and for PLA was 70% infill, high strain rate (80% / s). For the polymer blends, the sample / compressive test parameters were70% infill, high strain rate. The polymer blends studied were Ecovio® T2308 ((>75% PLA ± <25% PBAT), C7575 custom blend (75% PLA + 25% PBAT), C6040 custom blend (60% PLA + 40% PBAT), and Ecovio® F2224 (45% PLA + 55% PBAT). The results show that the custom blends exhibit material properties aligned with the properties of the commercial blends. In addition, the stress relaxation properties of the PLA + PBAT blend Ecovio F2341 at 25% infill, 50% infill and 70% infill were tested. The equilibrium modulus of Ecovio F2341 varies with 3D printed infill percentage, increasing significantly from 3 MPa at 25% infill to 7 MPa at 50% infill. There is no significant difference in equilibrium modulus between the 50% infill and 70% infill samples. The dynamic modulus of Ecovio F2341 increases with increased printed infill percentage, but with no difference in modulus for the same infill at different frequencies from 0.01 Hz to 10 Hz. The equilibrium modulus for 25% infill, 50% infill, and 70% infill samples is approximately 2 MPa, 5 MPa, and 7 MPa, respectively, for all measured frequencies. The dynamic mechanical analysis (DMA) phase angle for Ecovio F2341 increases with increased printed infill percentage and increases with increased frequency from 0.01 Hz to 10 Hz.

[0204] Additional mechanical testing of PBAT and PLA + PBAT polymer blends was performed. All samples were 70% infill and tested at a high strain rate (80% / s). The polymersDocket 88800730-000525tested were PBAT and PLA, and the polymer blends tested were Ecovio® F2341 (10% PLA + 90% PBAT), Ecovio® F2224 (45% PLA + 55% PBAT), C5545 custom blend (55% PLA + 45% PBAT), C6040 custom blend (60% PLA + 40% PBAT), C7525 custom blend (75% PLA + 25% PBAT), and Ecovio® T2308 ((>75% PLA + <25% PBAT). Properties evolve monotonically with composition, as expected. Results suggest that C5545 may be used for the “articular layer” material in the presently disclosed biologic living TKA joint design due to its suitable Young’s modulus, whereas C7525 may be used for the “bone” region. However, since the transition in properties between C5545 and C7525 is somewhat abrupt, as noted in C6040, some of these measurements of custom blends will be repeated to check reproducibility and iterate the blending process to achieve more reliable outcomes. Furthermore, in contrast to C6040, the yield stress of C5545 may be too low for the presently disclosed biologic living TKA joint design.

[0205] The design for the presently disclosed biologic living TKA joint to be used for finite element modeling (FEM) analysis were modified from CAD models of a knee implant. First, modularity was removed and smooth surfaces were created (FIG. 12A). Further modification involved introducing separate articular and bone scaffold domains for both the femoral and tibial components of the implant (FIG. 12B). FEM of the design for the presently disclosed biologic living TKA joint was performed using tensile material properties of solid and porous polymer blends. Initial FEM models were as follows: Model 1 - 50% porosity polylactic acid (PLA) bone domain, poly (1,4-butylene adipate-co-terephthalate) (PBAT) cartilage domain; Model 2 -Ecovio® T2308 (>75% PLA + <25% PBAT bone domain, PBAT cartilage domain. Load conditions were full extension, 10 body weights, representative of running forces. Peak stresses (same for both models) - bone domain - 9 MPa (tension), 21 MPa (compression), cartilage domain - 2 MPa (tension), 16 MPa (compression. The von Mises (effective stress of the Model 1 implant is shown in FIG. 12C, and the contact pressure of the Model 1 implant is shown in FIG. 12D. Using a safety factor of 3 shows that the bone domain scaffold has a tensile strength of > 27MPa and a compressive strength of >63 MPa, and the cartilage domain scaffold has a tensile strength of >6 MPa and a compressive strength of >48 MPa.

[0206] FEM on initial scans of commercial knee revealed non-congruent contact geometry between the femoral and tibial components. The FEM conditions were 10 body weights of load near full extension applied (simulating peak load during running), PLA-PBAT C7525 properties used for bone domain, PLA-PBAT C6040 properties used for cartilage domain. The FEM setupDocket 88800730-000525for the applied force of 10 body weights is shown in FIG. 13 A. The results showed high contact pressure (~27 MPa) and poor congruence (non-uniform contact area (FIG. 13B).

[0207] In silico plastic deformation was performed in the commercial knee scan models to create the presently disclosed biologic living TKA knee design based on post-deformation structure. Since the UHMWPE insert in the commercial knee may undergo plastic deformation during use, the presently disclosed biologic living TKA knee design does not require the same resilience to plastic deformation. The FEM conditions were 10 body weights of load near full extension applied (simulating peak load during running), and plasticity (elastic-perfectly plastic behavior) to model tibial “cartilage.” The results showed that plastic yielding occurs over most of the tibial articular layer, with the greatest plastic deformation occurring on the medial side (FIG. 14A andFIG. 14B).

[0208] FEM was performed on one embodiment of presently disclosed biologic living TKA joint (based on the plastically deformed commercial knee scan model). The FEM conditions were 10 body weights of load near full extension applied (simulating peak load during running), PLA-PBAT C7525 properties used for bone domain, PLA-PBAT C6040 properties used for cartilage domain. Contact pressure before altering tibial geometry is shown in FIG. 15 A, and contact pressure after altering tibial geometry is shown in FIG. 15B. The results showed reduced peak contact pressures and more congruent contact geometry. The peak decreased from ~27 MPa to ~14 MPa, and the new peak contact pressure is less than the yield stress of C6040 (~19 MPa).

[0209] To generate cartilage domain layer on the femoral component, a finite element mesh of the femoral component was created from the SolidWorks CAD model. A uniform pressure was applied on the articular surface using FEM. The deformed geometry was saved when the deformation approaches the desired average “cartilage” thickness. The deformed articular surface was extruded back to the original surface and this extruded domain was defined as the “articular layer.” The resulting model (FIG. 16A, FIG. 16B and FIG. 16C) has the requisite mesh needed for additional FEM.

[0210] To generate cartilage domain layer on the tibial component, a sectioning plane in SolidWorks was used to cut through the CAD model tibial component at desired average cartilage thickness and desired inclination (e.g., posterior) (FIG. 17A and FIG. 17B). A finite element mesh was created for each of the split domains, and a “tied contact” interface is used to simulate continuity of motion across that interface.Docket 88800730-000525

[0211] Studies were conducted using the Prusa-Slicer 3D printer software infill patterns versus custom CAD-created infill patterns. All samples used PLA with 70% infill (30% porosity). For the custom CAD model using a 0.8 mm nozzle, the yield strength was 44.8+ 1.4 MPa, Young’s modulus was 700 + 34 MPa, the toughness was 64.1 + 1.4 MPa, and the actual infill was 83% (17% porosity). For the custom CAD model using a 0.4 mm nozzle, the yield strength was 39.23M.4 MPa, Young’s modulus was 830 + 47 MPa, the toughness was 51.2 + 1.4 MPa, and the actual infill was 69.7% (30.3% porosity). For the Prusa-Slicer model using a 0.8 mm nozzle, the yield strength was 36.3 + 1.7 MPa, Young’s modulus was 560 + 41 MPa, the toughness was 45.23^1.7 MPa, and the actual infill was 73.4% (26.6% porosity). For the Prusa-Slicer model using a 0.8 mm nozzle and 100% infill, the yield strength was 95.3 + 10.3 MPa, Young’s modulus was 1150 + 260 MPa, the toughness was 69.3 + 2.8 MPa, and the actual infill was 995% (0.5% porosity). Results show that the custom CAD model with 0.8 mm nozzle produces significantly lower porosity (higher infill pattern, and thus higher compressive yield stress) than the Prusa-Slicer software. Therefore, custom-generated CAD patterns are not as reliable as those generated by the Prusa 3D printer software since they don’t reproduce the desired porosity. Using the 0.4 mm nozzle with custom CAD patterns produced poorer properties than the results using the Prusa-Slicer software. These results suggest it is better to use the Prusa 3D software to generate the desired porous patterns with desired infill percentage, and it is better to use the 0.4 mm nozzle than the 0.8 mm nozzle.

[0212] Failure testing was performed on one embodiment of the presently disclosed biologic living TKA joint. Iterating on testing jigs and protocols for failure testing and FEM validation was performed. One embodiment of the presently disclosed biologic living TKA joint was 3D printed using a specific design. 3D print matching loading jigs to mount on Instron testing machine at desired flexion angle. The 3D printed knee was loaded up to 10 body weights (e.g., 8000 N), and checked for failure from load-displacement response and micro-CT scanning. Pictures of the tibial loading jig, the femur loading jig, the tibial component and the femoral component before the test is shown in FIG. 18A, and after the test is shown in FIG. 18B. The results of the load displacement response showed a smooth upward curve as load and displacement increased. In this test 70% infill PLA was used for the entire implant geometry (no distinction between “cartilage” and “bone” regions). The loading jigs were printed at 100% infill to make them stiffer and stronger. The joint was loaded at 10 BW compressive load. The load-Docket 88800730-000525displacement response (and “before” and “after” photos) suggest no gross damage for this particular design and these particular material choices. The corresponding finite element analysis also predicts no failure under these conditions.

[0213] Mechanical testing of one embodiment of the presently disclosed biologic living TKA knee joint was performed. For this test the knee joint does not include “cartilage” and “bone” layers, and it does not employ the best choices of polymer blends for these two regions. The knee joint was 3D printed (FIG. 19A), was loaded to 10 body weights, and FEM was performed. The mechanical testing results showed no sudden drop in load-displacement response, suggesting no failure, and before / after micro-CT scans showed bi evidence of damage.

[0214] Custom polymer blends were prepared by melt blending 60% PLA and 40% EcoFlex® (PBAT) pellets. The pellets are fed into a 3D evo extruder for extrusion, and the filaments are collected (first melt blend). The first melt blended PLA + PBAT filaments are shredded, and the second melt blend is extruded and the filament is spooled. Alternatively, the pellet mixture of 60% PLA and 40% EcoFlex® (PBAT) pellets can be combined and mixed with chloroform at 15w / v%, the cast the polymer and evaporate chloroform into thin sheets for shredding, shred the polymer sheets into granulate, and extrude and spool the filament.

[0215] The current design for the NOVAJoint scaffold has four distinct 3D printed layers (FIG. 20): 1) 70% infill rectilinear coating layer of Ecovio T2308 to minimize friction; 2) 70% infill alternating (30° / 60°) triangular cartilage-seeding layer of Ecovio T2308; 3)100% infill rectilinear impermeable layer to prevent cell migration between domains; and 4) 70% infill alternating (30° / 60°) triangular bone-seeding layer of Ecovio T2308 + hydroxyapatite + 13-tricalcium phosphate.

[0216] While the overall process for designing the knee scaffolds has not changed (i.e., scan of commercial implant, smoothing process for initial CAD file, in silico deformation to minimize peak stresses, final design for 3D printing), an additional 3 mm slab has been added at the center of the back of the femoral implant to accommodate the four distinct layers throughout the entire implant (FIG. 28A and FIG. 28B). The updated scaffold design (FIG. 24A) also requires the design of a custom blade guide for surgical implantation (FIG. 33B and FIG. 33 C), which will produce an additional cut in the femur (FIG. 33A).

[0217] Different surface treatments were explored. Chondroitin sulfate (ChS) conjugation to Ecovio T23083D printed scaffolds was accomplished via grafting with NaOH functionalizationDocket 88800730-000525and drying overnight on day 1, and EDC-NHS treatment for 4 hours with agitation and finally submersion in ChS overnight on day 2. Toluidine blue staining confirmed the presence of ChS. The ChS conjugated scaffold has a reduced water contact angle and friction coefficient compared to a control Ecovio T2308 sample.

[0218] Ecovio T2308 scaffolds were coated with a commercial PTFE spray. Spray is clear and does not visibly alter the appearance of the scaffolds. The water contact angle after spraying increases slightly, but not significantly. The friction coefficient is reduced compared to a control Ecovio T2308 sample and a ChS-conjugated Ecovio T2308 sample.

[0219] CAD renderings of scanned traditional TKA implant femoral components were generated (FIG. 23). CAD renderings were then generated of femoral components of NOVAKnee scaffolds designed through in silico deformation (FIG. 24A), with exploded views of the sliced regions of the femoral component (FIG. 24B). Side views of CAD rendering of layered femoral components are shown in FIG. 28 A and FIG. 28B. CAD renderings of scanned traditional TKA implant tibial components were generated (FIG. 25). CAD renderings were then generated of tibial components of NOVAKnee scaffolds designed through in silico deformation (FIG. 26A), with exploded views of the sliced regions of the tibial component (FIG. 26B). Prior to testing NOVAJoint implants in cadaver knees, they were tested ex situ using an Instron testing system, which allows the application of 10 body weights (or more) to the articulating joint. The success of this experiment suggests that this approach can be used as the NOVAJoint design is refined.

[0220] A full 3D printed multilayer NOVAKnee construct was produced. The complete construct is shown in FIG. 27A, the femoral component is shown in FIG. 27B, and the tibial component is shown in FIG. 27C.

[0221] FEM analysis was performed to determine maximum force measurements for a tibial component of a commercial knee implant. The contact pressure (compression) was 52.4 MPa (FIG. 29A), and the maximum principal stress (tension) was 8.2 MPa (FIG. 29B). FEM analysis was also performed to determine maximum force measurements for the tibial component of the in silico deformation designed NOVAKnee. The contact pressure (compression) was 30.6 MPa (FIG. 30A), and the maximum principal stress (tension) was 13.0 MPa (FIG. 30B).

[0222] Representative overall dimensions of NOVAKnee sized to be similar to a size 9E of a commercially available conventional knee implants. The dimensions of the femoralDocket 88800730-000525component is shown in FIG. 31 A, FIG. 3 IB, FIG. 31C, FIG. 3 ID and FIG. 3 IE, and the dimensions of the tibial component is shown in FIG. 32A, FIG. 32B, FIG. 32C and FIG. 32D.

[0223] Surgical guides and jigs have been developed for NOVAJoint. A cut guide for the femoral component of NOVAKnee produces an additional cut in the femur to accommodate the central slab (FIG. 33A). A CAD rendering of the cut guide is shown in FIG. 33B (top view) and FIG. 33C (bottom view).

[0224] The process to generate cartilage filament is detailed below. The material used are PLA pellets, T2308 pellets and a filament spool. The dry pellets are added to a filament extruder, which is then cleaned and the data is exported. The process to generate bone filament is detailed below. The material used are PLA pellets, T2308 pellets, ceramic plus T2308 sheets and a filament spool. The ceramic plus T2308 sheets are shredded, generating dry pellets that are added to a filament extruder, which is then cleaned and the data is exported.

[0225] Cell seeded (human synovial MSCs) human NOVAKnee scaffolds were cultured for up to 28 days. Mechanical testing showed that properties were maintained over culture time. Force-displacement curves showed continuous curves (z.e., no discontinuity) indicating no failure of the implants over 3 repeated applied loads (8.5 mm / s testing speed, 700N, ~lx body weight) with 2-minute recovery time in between. Cells remained viable as noted by live-dead staining. As anticipated, mechanical behavior was unchanged with culture.

[0226] Cell seeded (human synovial MSCs) canine NOVAKnee were cultured for up to 28 days. Mechanical testing showed that properties were maintained over culture time. Forcedisplacement curves showed continuous curves (i.e., no discontinuity) indicating no failure of the implants over 3 repeated applied loads (8.5 mm / s testing speed, 600N, ~2.5x body weight, assuming a 24.5 kg dog at 45° flexion) with 2 minutes of recovery time in between. Cells remained viable as noted by live-dead staining. As anticipated, mechanical behavior was unchanged with culture.

[0227] Osteochondral constructs were examined in culture with and without loading. Osteochondral scaffolds (3.2 mm tall x 6 mm diameter, bilayer T2308 and 30% ceramic with impermeable layer) and iCPs in methacrylated hyaluronic acid (HAMA) hydrogel and iOsteo were cultured for up to 28 days with and without applied reciprocal shear loading (3 hrs / daily, 10 MPa). Cells remained viable (DNA) and produced cartilage ECM (GAG, glycosaminoglycans). TheseDocket 88800730-000525measures were not negatively affected by application of physiologic loading and provide in vitro support for the ability of these constructs to survive post-implantation in vivo.

[0228] EXAMPLE 3 - 3D-Printed Tissue Engineering Scaffolds

[0229] In regenerative medicine, fused deposition modeling (FDM) melts thermoplastic fdaments by a heating block, and then a nozzle head directs the extrusion of the melted filaments to deposit thin layers. There are wide ranges of biodegradable and biocompatible synthetic materials, or filaments, which can be printed. Synthetic polymers have been widely used as a class of biomaterials for both bone and cartilage replacement purposes. This is because their properties have been tuned in such a way to cover a broader range of elastic mechanical properties than natural biomaterials.

[0230] For instance, polylactic acid (PLA) possesses relatively high mechanical properties; but the material is also brittle, which limits its application for deformable load-bearing tissues. Mechanical testing on solid PLA dog bone-shaped samples was performed to test for static failure. The samples were printed on a PRUSA 3D-printer using 0°, 45° and 90° rasterization, including an edge to avoid potential stress risers, and tested on an Instron 5569A material testing system. The dog bone samples with 45° rasterization before and after testing were examined, and the results are summarized in Table 3, below:Table 3

[0231] Tensile tests were performed on 3D printed solid PLA using the same conditions (100% infdl, 0.4 mm nozzle height, 0.2 mm layer height, alternating rasterization) on different printers of the same model (Prusa Mark III). The average ultimate tensile stress, the average elasticDocket 88800730-000525modulus, and the average failure elongation were determined. The results showed variability that is statically significant, but not functionally significant.

[0232] Experimental testing of scaffolds in tension and compression demonstrated alternating infill pattern produces better material properties (including failure response) than unidirectional infill pattern, and the properties decrease nonlinearly with decreasing infill percentage (increasing porosity). 70% alternating infill is a good compromise for providing suitable pore size (~0.3 mm) and strength, compared to 50% and 25% infill. The alternating infill pattern at 70% represents an optimal pore structure for bone and cartilage layers.

[0233] Mechanical properties of 3D printed solid PLA and PCL was measured by uniaxial tensile tests. Solid PLA and PCL dog bone-shaped and elongated samples were tested. The average yield stress, average ultimate stress, and average elastic modulus are significantly higher in PLA compared to PCL when tested under tension. PLA has a yield stress of 42.13 MPa, ultimate stress of 42.82 MPa, and an elastic modulus of 1.42 GPa. PCL has a yield stress of 7.92 MPa, ultimate stress of 12.02, and elastic modulus of 0.15 GPa. PCL has a significantly higher average failure elongation than PLA under tension. PCL samples elongated to an average of 942.68% of their original size at failure. PLA samples elongated to an average of 5.17% of their original size at fracture.

[0234] Print settings were compared for their mechanical properties. Both prints used alternating infill (0° / 45o / 90o / 135°), 90 mm / s infill speed, 1 perimeter and 100% infill. The first print setting used 0.4 mm nozzle, 0.2 mm layer height, and the second print setting used 0.3 mm nozzle, 0.1 mm layer height. The effect of 3D printer nozzle diameter and layer height on PLA tensile properties was evaluated. 100% density alternating (0°, 45°, 90°, 135°) rectilinear infill samples were printed with either a 0.4 mm nozzle and 0.2 mm layer height (Group A) or a 0.3 mm nozzle and 0.1 mm layer height (Group B). There is a significantly different increase in average elongation %, average ultimate stress, and average elastic modulus for the samples in Group A compared to Group B.

[0235] Tensile tests were performed on elongated solid PCL samples. A representative stress-strain curve of PCL shows a linear increase in stress up to nearly 8 MPa at below 50% strain, then a near constant stress around 6.5 MPa during elongation from around 100% strain to 500% strain. Stress then increases nonlinearly up to nearly 11 MPa until failure at 800% strain. AnalysisDocket 88800730-000525of the PCL test specimen during testing show a distinct necking in the dog bone sample during the test and the sample did not fracture at the full elongation possible on the testing device.

[0236] To increase the toughness, a polymer composite can be made by adding plasticizers, rigid fillers and copolymers to improve the mechanical properties. Polycaprolactone (PCL), which is another FDA-approved biocompatible and bioresorbable polyester, can improve the toughness of brittle PLA. However, there are technical limitations for the use of PCL with PLA using FDM because the melting temperatures of PCL is much lower than PLA making melt blending of the two polymers incompatible for this type of fabrication, without burning the PCL. Given the unique deformability demands that are needed for cartilage and bone, unique polymer composites are needed to create a scaffold material using FDM that has low risk of brittle failure.

[0237] Building on the inventors extensive in vitro and in vivo experience with 3D printed (3DP) tissue engineering scaffolds, as well as the inventors understanding of tissue and joint biomechanics, new scaffolds are developed that permit 3D printing fabrication of the presently disclosed biologic living TKA joint components with deformable regions-specific mesh pore size and mechanical properties (stiffness and strengths) that permit immediate load bearing upon implantation. The use of flexible 3DP for biological applications including tissue engineering materials remains largely underutilized. Nevertheless, 3D printing of flexible polymers has shown promise in skin patches, drug delivery, and blood vessel modeling. In addition to endowing scaffolds with physiologically relevant deformability and energy absorption, flexible co-polymer reduce the potential of brittle failure due to mechanical loading and enhance the maneuverability or handling of tissues during implantation. The cartilage domain employs ceramic-free polymer blend that is more compliant and ductile, whereas the bone domain includes a composite of polymers containing ceramics to assist in osteoinduction. The addition of ceramics tends to make the polymers brittle, however the use of the ductile blend containing PLA and PBAT reduces this risk. The 3D printed scaffold allows for the manufacturing of custom, anatomically-shaped scaffolds.

[0238] 3D printing of porous osteochondral scaffolds with various polymers and polymer blends was performed, and the compressive equilibrium modulus was measured. Porous biomaterials were created from PLA, Ecoflex® (100% PBAT) and PCL using 3DP. Sample details: 15mm diameter x 6 mm height;. 25% infill (75% void space); alternating infill pattern (90, 45, 135, 0 degrees). Printed using Prusa MK4 with no perimeter layer. Compressive dynamicDocket 88800730-000525mechanical analysis (DMA) of the porous scaffolds was performed. Compressive dynamic mechanical analysis (DMA) of PLA, PCL, and Ecoflex (PBAT) with 75% porosity (25% infdl) resulted in a modulus of approximately 20 MPa, 10 MPa, and 1 MPa for each material, respectively, across a frequency range of 0.1 - 10 Hz. The phase angle of all three materials is approximately equal increasing from approximately 1 degree at 0.1 Hz to approximately 5 degrees at 1 Hz, 20 degrees at 5 Hz, and 40 degrees at 10 Hz. Porous cylinders were tested under unconfmed compression stress relaxation test, with 2% strain applied at 0.0217 mm / s (based on D695 ASTM standard) after a 5N preload was used. Within each porosity (% infill), materials have significantly different equilibrium moduli: 25% Ecoflex® (2.2MPa) < PCL (6.8 MPa) < PLA (19.9 MPa); 50% Ecoflex® < PCL < PLA; 70% Ecoflex® <PCL < PLA. The equilibrium modulus for Ecoflex, PCL, and PLA increases with increased infill for all three materials. The equilibrium modulus for each sample is in Table 4 below.Table 4

[0239] The porous polymers exhibit porosity dependent (% infill) equilibrium compressive moduli. N=6 per group. An ANOVA was performed to compare equilibrium modulus as a function of % infill with Tukey post hoc test. Significance shown for p<0.05. Porous PCL exhibits frequency, but not porosity, dependent dynamic compressive moduli. Porous PCL exhibits frequency and infill percentage dependent dynamic compressive moduli. Porous EcoFlex® exhibits frequency and infill percentage dependent dynamic compressive moduli.

[0240] The porous polymers exhibit a range of dynamic compressive moduli that are within the target dynamic properties for cartilage domain (15 MPa - 60 MPa). Porous cylinders of PLA, PCL and EcoFlex® were tested under unconfined dynamic mechanical testing with N=6 for each polymer. Testing configuration: 0.5% strain applied at 0.01, 0.1, 0.5, 1 and 10Hz.Docket 88800730-000525Samples with 75% porosity exhibit a range of compressive dynamic modulus with PLA > PCL > Ecoflex®.

[0241] The 3D printed mesh constructs (25% gyroid infill) are subjected to degradation characterization. The samples were submerged in water (pH 7.4) at 37C with continuous agitation. The dry and wet weight of the samples are collected at day 0 and study endpoints (12 months) to calculate the percent mass loss. Degradation pilot study on N=3 per group and time point. ANOVA was performed to compare differences between groups and demonstrated no significant difference in % mass loss at 1 month time point, a small, but statistically significant increase in remaining mass of PLA at 2 months. EcoFlex® and Ecovio® show minimal mass loss. In vitro hydrolytic degradation studies of PLA, Ecovio-75% PLA, Ecovio-45% PLA, Ecovio-10% PLA, Ecoflex, and PCL with 25% gyroid infill density show that all of the materials except for Ecovio-10% PLA have minimal (< 0.5%) mass loss over 5 months. Ecovio 10%-PLA has approximately 3% mass loss at 4 months. The porous scaffolds are stable in vitro, with <0.2% mass loss at 3-month timepoint. EcoFlex® ( 10% infill) showed the most obvious evidence of degradation though still very small (<1%).

[0242] Porous osteochondral scaffolds with 600 pm fiber diameter were printed using an Ultimaker S3 printer and a Prusa i3 MK3S+ printer, and the results were examined. The scaffolds printed with the Prusa i3 MK3S+ printer yielded better morphology / print fidelity (Table 5).Table 5

[0243] Further testing was performed using porous osteochondral scaffolds with 600 pm, 500 pm, and 400 pm fiber diameter printed using a Prusa i3 MK3S+ printer. Fidelity in controlDocket 88800730-000525of resulting fiber spacing improved with decreasing fiber diameter, where smaller differences between nominal and actual fiber spacing was observed (Table 6).Table 6

[0244] Candidate polymers poly (1,4-butylene adipate-co-terephthalate) (PBAT; Ecoflex®) or PBAT-PLA (Ecovio®), which have desirable elastomeric properties (easily stretched with large, recoverable strain, and mimics the mechanical properties of Flexifil), and are 3D printable (extrudable / melt flow index, and PLA based, which is relevant to biological implantation), were studied, using Flex-Fil as a standard (Table 7). An increase in the percentage of PLA in Ecovio results in an increased tensile modulus and decreased elongation at failure. Tensile testing results for various compositions are in Table 8, below.Docket 88800730-000525Table 7Table 8Docket 88800730-000525

[0245] The inventors implemented melt extrusion of two biodegradable polymers, polylactic acid (PLA) with the polymer PBAT. Melt blending of ductile PBAT with brittle PLA improves the toughness of composite and reduces the risk of brittle failure of the implant. Both materials are also biodegradable, biocompatible, and are compatible with each other for melt blending due to a similar melting temperature profile. PLA-PBAT (80 / 20) blends were extruded into polymer filaments, spooled and used for 3D printing of cartilage and bone scaffolds (implant). The internal structure of the scaffolds was optimized to increase mechanical integrity while permitting matrix deposition in cartilage domain and bone ingrowth in boney domain. A porosity of 30% (70% infill density) was identified as optimal. Fiber orientation was also evaluated to enhance mechanical integrity. The inventors found that sequential layers deposited with a rectilinear fiber pattern that alternates by 45° in consecutive layers to provide the greatest resistance for failure in tension and compression.

[0246] In certain embodiments, the temperature for proper extrusion of filaments containing PLA and / or PBAT is 170°C to 200°C and for 3D printing of scaffolds containing PLA and / or PBAT in layer by layer process is 200°C to 220°C. In certain embodiments, the printing of the hydrogel for the scaffolds is at room temperature, or about 40°C.

[0247] Using the PRUSA software, an example of a two-layer tibial implant component was produced (FIG. 34), with parts printed in different materials imported as separate files, assigned a material, and then printed together. The Prusa 3D printer was upgraded to include a multi-material printhead to allow for 3D printing of multilayer scaffolds with different materials for cartilage and bone.

[0248] Anatomical knee 3D printing led to challenges with the smoothness of the articular surfaces. Complex non-planar shapes (FIG. 35 A) require the use of support materials (yellow portion of FIG. 35 A) during 3D printing. Physical removal of the support surfaces led to microabrasions and a roughened surface (FIG. 35B). To circumvent this, a water dissolvable material support material was used (PVA) (FIG. 35C). The resultant process for 3D printed multilayer osteochondral scaffolds for the tibial and femoral component is shown in FIG. 35D.

[0249] Representative pCT of 3D printed porous scaffold microarchitecture (triangular infill) is shown in FIG. 36A (side pores), FIG. 36B (fiber width), FIG. 36C (pore height) and FIG. 36D (pore width).Docket 88800730-000525

[0250] Mechanical properties of acellular bone and cartilage scaffolds, before and after sterilization were investigated. Ecovio T2308 compressive properties were evaluated on cylindrical samples and tensile testing on dog bone samples pre- and post-sterilization using an Instron mechanical test system. The measured yield stresses for both compression and tension are below the peak tensile and compressive stresses as determined through finite element (FE) models of NOVAKnee-T designs. No change in tensile or compressive mechanical properties of Ecovio T2308 after ethylene oxide sterilization was noted for the cartilage scaffold.

[0251] In vitro hydrolytic degradation of cartilage domain scaffold was investigated. Degraded scaffolds were evaluated for mass loss or loss in mechanical properties. SEM of T2308 (25% gyroid infill) before and after 11 months of degradation testing in distilled water was examined. Though there has been no appreciable mass loss in 12 months of hydrolytic degradation in vitro at 37°C the initiation of degradation is observable at the micro- and nanoscale. Ecovio T2308 scaffolds (25% gyroid) incubated in distilled water for up to 12 months maintains their integrity, with about 25% loss in dynamic modulus.

[0252] In vitro hydrolytic degradation of bone domain scaffold was investigated. (ASTM Standard - F1635 - 16), scaffolds of T2308+HT 70% infill alternating triangular pattern were immersed in PBS at 37°C for three months. Results indicated that T2308+HT degrades at a faster rate than T2308 alone. Both scaffolds have less than 2% mass loss after six months and remain fully intact at the macroscale. Though there has been no appreciable mass loss in three months of hydrolytic degradation in vitro at 37°C, surface changes are observable at the micro- and nanoscale via SEM.

[0253] Fatigue testing of scaffold material was investigated. Fatigue testing of Ecovio-T2308 with 70% alternating triangular pattern (30° / 60°) tested at 4 MPa and 3000 RPM. Desired fatigue failure criterion is at least 1 M cycles, based on an approximation of the number of steps older adults take annually. In initial studies, six samples met this criteria, with five of six samples showed no failure up to 5 million cycles and one sample failing at 1.05 million cycles.

[0254] In vitro cytotoxicity of scaffold (ISO- 10993 -5) was investigated. For the cytotoxicity extraction study, ethylene oxide sterilized Ecovio T2308 samples were used. T2308 scaffolds were submerged in media for 3 days at 37°C to allow any potential factors to be extracted into the media. Cells (L-929 fibroblasts) were then incubated in extracted media at multiple dilution levels (1:1, 1:2) for 24 hours to determine cytotoxicity of materials. For the direct contactDocket 88800730-000525study, L929 fibroblast cells were incubated with sterilized scaffolds for 24 hours. In both experiments, effects of Ecovio T2308 were compared to rubber latex as a positive control and LDPE or HDPE as a negative control of cytotoxicity. As expected, rubber latex led to low viability in extracted media as well as in direct contact (<70%). Similarly, LDPE was found to be minimally cytotoxic to cells. Cells exposed to Ecovio T2308 extracted media or in direct contact with scaffolds yielded high level of cell viability (>70%).

[0255] In vivo immunogenicity of NOVAKnee scaffolds was investigated. 3DP multilayer osteochondral scaffolds (T2308 cartilage layer, an impermeable interface, and T308HT boney layer) were subcutaneously implanted in wild-type, immunocompetent rats. At 6 weeks post implantation, the scaffolds were resected and processed for histological analysis to evaluate potential of a sustained foreign body or immune response, and for evidence of matrix deposition or angiogenesis. Minimal reactivity was found at 6 weeks post-implantation, using ISO 10993-6 histological scoring system. Thin fibrovascular tissue layer mostly surrounded the implant and extended into the center. Also, minimal evidence of mononuclear cells was observed. Evidence of neovascularization observed in the boney layer of all samples. Histology of acellular scaffold 6-weeks post-implantation was examined.

[0256] Mechanical assessment of in vivo NOVAKnee scaffolds was investigated. 3DP multilayer osteochondral scaffolds (T2308 cartilage layer, an impermeable interface, and T308HT boney layer) were subcutaneously implanted in wild-type, immunocompetent rats. At 6 weeks post implantation, the scaffolds were resected and were evaluated for change in scaffold shape and degradation (mass loss). Scaffolds were also tested mechanically to evaluate changes in compressive properties post implantation. There was a small mass loss in vivo (<5%) with no major change in scaffold morphology. Mechanical testing showed similar yield strength compared to pre-implantation (day 0) controls, with slight decrease in compressive modulus and ultimate strength, and slight increase in compressive toughness. The stability of the compressive yield strength with in vivo implantation meets the compressive design criteria for NOVAKnee-T.

[0257] EXAMPLE 4 - Cell Sourcing

[0258] The presently disclosed biologic living TKA joint delivers articular chondrocyte and osteogenic lineages that form cartilage and bone, respectively, post-implantation. Using the most recent advances in stem cell biology and cell manufacturing, human induced pluripotent stem cell (iPSC) derivatives serve as the allogenic cell source for the allogenic biologic living TKADocket 88800730-000525joint. Cell differentiation protocols to generate optimal cells to bank are performed using cutting edge robotic platforms that provide population-scale, high-throughput cell culture methods. Autologous cells utilize adipose-derived stem cells (ASCs) obtained from minimally invasive lipid aspirates. However, differentiation potential for autologous ASCs may be variable and donorspecific.

[0259] Single cell RNA sequencing (sc-seq) are used to identify metrics that predict the potency of ASC differentiation and cartilage / bone production, which are used to determine whether compensatory measures are needed to mitigate low performing patient cells. These techniques identify good vs poor donors by correlating the sc-seq datasets with in vitro differentiation potential. Surface markers are identified that can be used to rapidly screen donor ASCs to determine whether cells with good differentiation potential are present in sufficient numbers.

[0260] Protocols for differentiations are based on prior published work (Adkar, etal., Stem Cells. 37:65-76, 2019. Loh, et al., Cell. 166(2):451-467, 2016. Kawai, et al., Nat Biomed Eng.3(7):558-570, 2019. Estes, etal., Nat. Protoc. 5:1294-1311, 2010; Jaiswal etal., J. Cell Biochem.64:295-312, 1997; Briggs etal., J. Biomed. Mater. Res. A 91:975-984, 2009).

[0261] Collected, isolated and expanded ASCs from 11 donors. Example images of ASCs from six donors were examined, which demonstrates the range in initial cell numbers. All cells exhibited good morphology at passage 0, and were banked at passage 0. Additional observations on banked ASC lines: 7 female donors, 4 male donors: 6 ASC lines were isolated the same day as the lipoaspirate collection, 5 were isolated the next day (not exceeding 24 hours - no obvious differences were observed between same and next day isolations); average time to confluency (freeze) from isolation was 14.4 + 4.5 days (n = 10); average number of cells at time of freezing 13.5 + 5 million cells (n = 10); roughly 100 ml of aspirate was collected for each patient, but the fatty portion varied somewhat between patients (9 / 11 donors had stromal vascular fraction (SVF) of 75-100 ml, 2 / 11 donors had SVF of 50-75 ml).

[0262] The ASCs from the 11 donors were characterized for time to freeze and cells at freezing based on gender, age and body mass index (BMI). No initial differences or correlations were observed between age and BMI as a factor in the time for isolated cells to reach confluency to freeze or with number of cells attained at time to freeze.Docket 88800730-000525

[0263] Naive hiPSCs were cultured and expanded. The WTC11 line was purified and expanded for differentiation studies. Two days post-thaw the cells showed improved post-thaw survival using hESC-qualified Matrigel™. The culture included a mix of differentiating and pluripotent colonies. A single colony was picked to purify cultures. After three days (single colony expansion) the hiPSCs showed characteristic morphology of hiPSCs.

[0264] ASCs were isolated and characterized. Flow cytometry show majority of isolated cells are ASCs (passage 1). Flow cytometry was conducted on ASCs from two donor patients to evaluate surface marker expression of ASCs after 1 week. For both patients, >59% of cells are positive for markers CD90, CD105, and CD73, while <2% of cells are positive for markers CD45, CD34, CD14, and CD19.

[0265] The tumorgenicity potential in iPSC derived osteogenic (iOsteo) cells was investigated. Thaw and expansion of iOsteo cells depletes non-differentiated iPSCs-the TRA-1-60A population that remain pluripotent and that may form teratomas in vivo. The use of PluriSin is also effective in killing pluripotent hiPSCs in 24 - 48 hours without killing iOsteo cells, and can be used if further reduction in persisting pluripotent populations is required.

[0266] In vitro minimization of pluripotency was investigated. iCP cell populations derived from GMP clone lines (PO) exhibit pluripotency markers (TRA160, SSEA4, CD30) in less than 0.05% of cells (Table 9). iOsteo cell populations derived from GMP clone lines (P0) exhibit pluripotency markers in less than 4% of cells after initial differentiation, but freezing, thawing, and expanding (Pl) pluripotent populations are further reduced to -0.7% (Table 10).Table 9Table 10Docket 88800730-000525

[0267] Thaw and expansion of iOsteo cells prior to experiment depleted TRA-1-60A population, therefore PluriSlnl treatment will not be required (no effect when TRA-1-60 levels are low). Differentiated iOsteo cells had high levels of TRA-1-60+ cells prior to freeze, thaw, and expansion. Flow cytometry detects very few TRA-1-60+ cells in iPSC-derived chondroprogenitors (iCPs). The values of pluripotent percent parent for iPSC, ASC, iOsteo WTC11, iOsteo SKIPS and iOsteo SK003.2 is shown below in Table 11.Table 11

[0268] PluriSin 1 is a stearoyl-CoA desaturase 1 (SCD1) inhibitor used to selectively eliminate undifferentiated human pluripotent stem cells (hPSCs) from culture; it induces ER stress, attenuates protein synthesis and induces apoptosis in hPSCs. Prevents teratoma formation in immunocompromised mice. iPSCs treated with Pluri Sin show signs of cell death within 24 hours and all cells killed by 48 hours. No naive hiPSCs made it to flow cytometry after PluriSlnl treatment. Morphology and viability of naive ASCs and iOsteo cells were good with PluriSlnl treatment. PluriSin could be used if further reduction in pluripotent populations is required.

[0269] Formation of osteogenic phenotype in vitro was investigated by measuring gene expression of RUNX2 and COL1 Al . Differentiation protocols were carried out on four different iPSC lines (research-grade SKIPS; GMP clones D, E, G) at 10, 12, and 14 days. Extending timeline from 10 days to 12 days further improved osteogenesis, but no benefit observed past 12 days. All GMP clones outperform the research grade line at 12 days. The results of Alizarin red staining of iOsteo cells from GMP clones D, E, and G after 10-day differentiation protocol was evaluated. GMP Clone E was selected for clinical study (best option based on combined results of chondrogenesis and osteogenesis).Docket 88800730-000525

[0270] Maintenance of osteogenic phenotype in vitro was investigated by measuring gene expression of RUNX2 and COL1A1 iOsteo cells that were frozen and thawed, and that were expanded in osteoinductive medium for one passage. Osteogenic genes continued to increase after one passage.

[0271] Reduced calcium staining with 1 mM beta-glycerophosphate (bGP) compared to 10 mM bGP during iOsteo differentiation (10 days) was seen, with only slight reductions in osteogenic gene expression. Gene expression of BGLAP, COL1A1 and RUNX2 was measured, with the 1 mM samples showing higher BGLAP expression, but lower COL1A1 and RUNX2 expression compared to the 10 mM samples.

[0272] Optimization of hiPSC osteoblast differentiation was investigated. The cells were treated with Accutase forlO or 20 minutes, with TrypLE for 10 minutes, trypsin for 2 minutes then Cis IV for 10 minutes, and Cis IV for 30 minutes. It was found that 15 min of Accutase treatment combined with aspiration led to total dissociation of iOsteo with 75% viability.

[0273] Osteoblast differentiation of hiPSCs was determined at 7 and 10 days with or without RA. Both osteogenic groups (with and without RA) showed successful staining of bone markers COL1 Al and BGLAP at day 10 of differentiation. A slight improvement of staining with RA treatment at Day 7 was seen. Successful osteogenesis of hiPSCs in was accomplished in 7-10 days. RA treatment significantly promoted COL1A1 expression. Osteocalcin and RUNX2 were only detected in the D10 group. There was no significant difference between D7+RA a d D10+RA.

[0274] A 21 -day chondrogenesis protocol was found to be acceptable for all NYSCF clones. iCPs collected at 21 days showed good expression of cartilage marker, and CollOal was not detected for any sample.

[0275] Osteochondral scaffolds in loading bioreactor were evaluated. Cell seeded (iCPs in HAMA hydrogel, iOsteo) cylindrical osteochondral constructs (porous T2308 / impermeable T2308 / porous T2308+HT) were cultured for up to 28 days with or without an externally applied load (10 MPa per construct, reciprocal shear loading, 3 hours per day). Cells remained viable (DNA) and produced cartilage ECM (GAG, glycosaminoglycans). These measures were not negatively affected by application of physiologic loading and provide in vitro support for the ability of these constructs to survive post-implantation in vivo. Alcian blue staining showed robust chondrogenesis at 3 weeks in iCPs from GMP clones D, E, and G.Docket 88800730-000525

[0276] HAMA hydrogel showed better structural stability with iCPs than FibGen:HA gel. iCP single cell suspensions in HAMA hydrogel and FibGen:HA gel were prepared. Unlike the FibGen:HA gel with iCP single cell suspension, which resulted in full degradation in a short time, HAMA maintained the original structure in chondrogenic media over 1 week.

[0277] Chondrogenic differentiation at 12 and 21 days for SKIPS hiPSC line was investigated. COL10A1 was not detected in iCPs at either 12 or 21 days. D21 iCPs have similar or significant increase of all chondrogenic markers compared to D12. Chondrogenic differentiation was investigated for three hiPSC lines (WTC11, SKIPS and SK003.2), and AC AN and COL2A1 expression was measured. hiPSC chondrogenesis is variable across different hiPSC lines - WTC11 line seems less effective than SKIPS or SK003.2. D21 iCPs have consistently more Alcian Blue staining, and therefore 21 days of differentiation will be used. SK003.2 iCPs showed the best response, so these cells are differentiated for SCID mouse studies. ACAN was only detected in SKIPS chondroprogenitors, iPSC-derived chondroprogenitors from all three lines have significant higher expression of COL2A1 expression, and SK003.2 somite has slightly detectable level of COL2A1 RNA. COL10A1 not detected for all samples.

[0278] Optimization of hiPSC chondrocyte differentiation was investigated. The cells were treated with Accutase for 10 or 20 minutes, with TiypLE for 10 minutes, trypsin for 10 or 15 minutes, and Cis IV for 30 minutes. High viability with 10 minutes Trypsin digestion (over 90%) was observed. While other digestion methods such as Accutase showed lower viability, there is some contribution of artifact because there were more cell clusters with this method which were detected as ‘dead’ cells even though they were not dead. Trypsin digest resulted in uniform single cell suspension. A remarkable number of cells can be recovered from a single well of a 12 well plate.

[0279] Immunogenicity evaluation of differentiated iPSCs was investigated. Cells (50K per well) were treated with mitomycin C to stop proliferation and then co-cultured with PBMCs for 7 days. Proliferation of PBMCs (calculated as stimulation index) indicate positive immune response (compared to the positive controls Allo and PHA). Initial immunogenicity results with SKIPS research line show no reaction with iOsteo cells and reaction against only one PBMC donor for iCPs (cells were tested against 3 donors). Flow cytometry of immunogenic markers MHCI, MHCII, CD40, CD80, and CD86 were either generally comparable to naive ASCs, or had reduced expression compared to naive ASCs, used as a control as minimally reactive cell type.Docket 88800730-000525

[0280] A canine iPSC line was derived. The line is being cultured out, and will be used in in vitro studies on cartilage and bone differentiation, and in vivo large animal studies.

[0281] In vivo assessment of osteogenic stability was investigated. iOsteo cells were loaded on T2308-HT scaffolds and implanted in SCID mice for 3 weeks. COLlal gene expression was reduced with in vivo implantation compared to starting iOsteo expression level; however, expression of RUNX2 (a specific transcription factor for bone) was 10-fold upregulated, suggesting continued osteogenic differentiation in vivo. Alizarin red staining to detect mineralization shows increased staining of tissues at the edges of the scaffold. MicroCT showed several bright spots that may be new bone or signal from the ceramic.

[0282] In vivo SCID mouse studies were conducted using iChondro pellets (instead of single cells). After 3 weeks of subcutaneous implantation, Alcian blue staining shows GAG-matrix forming around pellets, and Alizarin red staining shows no mineralization. In vivo chondrogenesis of iChondrocyte pellets in different sizes delivered via HAMA hydrogel was evaluated. Three weeks in vivo implantation showed cell pellets structures remained. Cartilaginous matrix was formed per Alcian Blue staining in and surrounding iChondrocyte pellets, but not much hydrogel degradation was observed. Single-cell iPSC-derived chondroprogenitors (iCPs) in HAMA show more robust chondrogenesis in vivo than in vitro. Micro-CT (pCT) of iCPs in HAMA harvested after 3 weeks in vivo in SCID mice was examined. No calcifications were detected at 3 weeks. Any bright spots were generally 15-52 mg HA / ccm while calcification is generally defined as >200 mg HA / ccm. Gene expression of Sox9 an ACAN for iCPs after 3 weeks in vivo (compared to bovine in vivo samples) was also performed.

[0283] iPSC-derived chondrogenic (iChondro) pellet size can be controlled consistently by cell number. Pellet size is consistent across groups (OAK, IK, 10K and 30K cells / pellet). All group diameters are significantly different from each other (p < 0.0001). Micro-CT of iChondro pellets encapsulated in HAMA gels at -50-60 pellets per gel and implanted in SCID mice for 3 weeks was examined. For the IK pellet size, 3 bright spots detected between 66.6 - 156.6 mg HA / cm, for the 10K pellet size, 2 bright spots detected between 121.5 - 179.5 mg HA / cm, and for the 30K pellet size, 3 bright spots detected between 70.7 - 717.3 mg HA / cm, with two of these above the 200 mg HA / cm threshold for calcification. No bright spots detected for 0.5K pellet size. Expression of ACAN, COL2A1 and COL10A1 was examined, and chondrogenic gene expression is generally maintained or higher for all in vivo pellets compared to controls, except OAK pellets.Docket 88800730-000525

[0284] A description of ASC isolation and cell expansion is shown below in Table 12.Table 12

[0285] Formation of osteogenic phenotype in vitro was investigated. Alizarin red staining shows the differentiation process for osteoblasts from ASCs (ASC-OBs) occurs beginning at 2 weeks with the current protocol. ASC-OBs were cultured on T2308+HT composite scaffolds in osteogenic media for 14 days. Osteocalcin production was quantified using ELISA. A > 60% increase in osteocalcin production was observed for T2308+HT scaffolds compared to T2308 alone, showing T2308+HT promotes osteogenesis. Alizarin red staining also provides qualitative evidence that ASC-OBs on T2308+HT scaffolds result in an increase in mineralization.

[0286] Formation of chondrogenic phenotype in vitro was investigated. Under pellet culture, upregulation of chondrogenic genes (ACAN and SOX9) is observed consistently across ASC samples from multiple patients after 2 weeks of differentiation in chondrogenic media. Robust and consistent Alcian Blue staining can also be observed for ASC micropellets (10K cells / pellet) cultured for 3 weeks in chondrogenic media.

[0287] Novel markers were identified and optimization strategies were developed to improve ASC differentiation potential using scRNA-seq. Single-cell RNA sequencing (scRNA-seq) was completed for ASCs collected from 11 patients to address donor variability. The scRNA-seq was performed at passage 1 to ensure that the cell populations sequenced were primarily ASCs. The goal is to first determine whether there are predictive cell markers or distinctive cell populations associated with donors that have high orDocket 88800730-000525nonhypertrophic chondrogenic / osteogenic potential. This information can be used to enrich cells to improve differentiations (by FAC sorting for example). The second goal is to identify potential signaling pathways that may be over-expressed in poor donors that may need to be suppressed using inhibitors or antagonists during differentiation to better push cells toward desired phenotypes.

[0288] A knee plot of all samples indicates high quality control. UMAP (Uniform Manifold Approximation and Projection) plots are visualizations that reduce complex, highdimensional data (like gene expression from thousands of genes) into a 2D or 3D map, revealing underlying patterns, clusters, and relationships. For example UMAP1 and UMAP2 can be used as new axes to simplify the plotting of the results. Clusters represent groups of cells with different patterns of transcriptome expression. Thus, each cluster represents a different group of cell types or same type of cell but with a different phenotype or function. The features that differ between the different clusters are the expression of many genes that define the identify of each cluster. Three primary clusters were identified from the initial analyses of UMAP visualizations of all 11 donors. Cells from most donors have cells in each cluster, but in some cases, cells skewed towards one or two clusters. Proliferating cells were mostly identified in Clusters #2 and #3. The expected MSC markers are expressed by all donors to largely equal extents, with no specific chondrogenic nor osteogenic differentiation genes expressed (except Coital which is a broad marker expressed by many stromal cell types, so not a marker helpful specific to osteogenesis though frequently used).

[0289] In previous differentiation studies with Alcian blue staining, there was no noticeable difference between donors for chondrogenesis at 3 weeks. However, there has been notable variation in osteogenesis between donors after 3 weeks of differentiation, based on Alizarin red staining. As such, representative donors were analyzed based on their osteogenic potential using Differentially Expressed Genes (DEG) testing, and a series of genes were identified that are upregulated or downregulated in great / medium donors compared to poor donors (Table 15).Table 15Docket 88800730-000525

[0290] EXAMPLE 5 - Chondroinduction

[0291] Currently available bioprintable hydrogels have limited mechanical properties, chondrogenic potential, and degradation rate. Fibrin-based hydrogels have been widely used for tissue engineering of musculoskeletal tissues, including cartilage. However, its mechanical properties and fast degradation, especially inside synovial joint, are remaining challenges.

[0292] An innovative combination hydrogel is incorporated into the interstices of the scaffold mesh and serve both a biological and mechanical function in the presently disclosed biologic living TKA joint. As a carrier for cells and chondro-inductive factors, the infused novel hydrogel creates an environmental niche to promote cartilage regeneration. The presently disclosed biologic living TKA joint employs a bioprinter that uses mechanical extrusion where hydrogel and cells or biofactors are combined to form a slurry and then plotted to manufacture the implant. This hydrogel is also designed to recapitulate the natural load-bearing mechanisms of articular cartilage: a high water content that supports interstitial fluid pressurization (accounting for more than 90% of the applied load and defining the low friction coefficient of cartilage) and to promote boundary lubrication via self-assembly of lubricin-coated articular surfaces upon immediate contact with synovial fluids. Optimal hydrogel compositions are established in a combination of heparin-conjugation (self-assembly of frictionless articular surface) and GAG-mimetic cellulose, which are chondroinductive and anti-inflammatory, and genipin crosslinking (initial strength and delayed degradation). Factors to modulate the pro-inflammatory OA environment can also be incorporated.Docket 88800730-000525

[0293] The inventors implemented genipin crosslinking of fibrin gel (FibGen) that significantly delayed degradation. In addition, optimized control of genipin concentrations resulted in significantly improved mechanical properties suitable for cartilage coating layer supporting cartilage formation in the presently disclosed biologic living TKA joint. The experimental data demonstrated the compressive modulus at a higher order of other hydrogel, and a high cell viability with chondrocytes and mesenchymal cells. To further enhance the chondrogenic potential, GAG-mimetics are being studied. The hydrogel composition will have FibGen, GAG-mimetics, and human derived chondrogenic cells.

[0294] Initial tests using the BioAssemblyBot 400 bioprinter to fabricate bone composite scaffolds of PCL was performed. The print parameters were 110°C and 100 psi.

[0295] However, the dual nozzle printing unit equipped in the BioAssemblyBot 400 system (Advanced Solutions, Louisville, KY), has significant limitations in printing FibGen. As the two solutions (Fibrinogen and Thrombin + Genipin) being fed are mixed inside the T-shaped channels, the hydrogel gets clotted in the conduit. In addition, hydrogel materials are dispensed by air-pressure which suffers from a limited accuracy in controlling printing rate.

[0296] To resolve the issues, the inventors modified the syringe holder to be equipped with a custom-designed, Y-shaped nozzle (FIG. 37), which allows a mixture of two materials right before being dispensed. The dispensing unit was also modified to hold two syringes and to apply pressure controlled by a stepping motor, allowing micro-precise control in printing feed rate. After a series of modification and calibration of the prototype (FIG. 38A), the final version of modified dual printing unit (FIG. 38B, FIG. 38C and FIG. 38D) was established, which resulted in homogenous control of hydrogel printing thickness and infiltration depth into the 3D-printed polymeric scaffolds. Two different types of Y-shaped channel nozzles were tested: Screw-type (FIG. 39A) and two-channel exit (FIG. 39B). Screw-type improved the printability, but the printed hydrogel showed a bulge-up structure. The two-channel exit model exhibited the superior outcome in printing FibGen hydrogel. The bioprinting of FibGen hydrogel onto 3D scaffolds was evaluated at different nozzle heights and motor speeds to fine tune control of the hydrogel thickness on the scaffold surface. These studies showed the hydrogel thickness was well-controlled by the steps / second speed of the stepper motor, with the FibGen thickness decreasing with a decreased number of steps / second. Changing the nozzle height from 0.5 mm above the scaffold surface to 1 mm above the scaffold surface resulted in a more even thin layer of hydrogel (based on visualDocket 88800730-000525inspection). The inventors optimized a number of printing parameters, including feeding rate, printing speed (movement of printing head), and surface treatment on the solid scaffold for cartilage, resulting in a controllable, homogenous hydrogel coating and infiltration depth.

[0297] Bioprinting of cartilage hydrogel was performed. Heparin-conjugated FibGen was prepared to promote adhesion of lubricin to the hydrogel. Hep-FibGen was tested for compression and compared with FibGen. Stress-strain curves show similar features between Hep-FibGen and FibGen. Hep-FibGen showed higher modulus at Zone 1 but no difference at Zone 2 and Zone 3. These data indicate the sufficient mechanical properties of Hep-FibGen.

[0298] Feeding rate optimization was performed for the homogenous layering of hydrogel on top of scaffolds. By controlling the feeding rate, controlled by the stepping motor spinning speed, the thickness of the bioprinted FibGen 100 was controlled.

[0299] Testing of bioprinting hydrogel on top of 3D printed solid scaffolds was conducted at a nozzle height of 0.5 mm and 1 mm. By adjusting the initial nozzle height formation of a thin layer of hydrogel (FibGenlOO) on 3D printed PCL scaffolds with 300 pm pores was observed.

[0300] Compression tests were conducted for 3D printed hydrogel with multiple compositions of fibrinogen, thrombin and genipin (fibrinogen (20-100 mg / ml) + thrombin (20100 mg / ml) and 0.5-2.5 mg / ml genipin). The compressive modulus and maximum strain were measured for FibGen concentrations ranging from 20 to 100 mg / mL fibrinogen. The compressive modulus generally increased with increasing concentration from 20 to 80, while being approximately 70 MPa for both FibGen80 and FibGenlOO. The maximum strain was approximately 0.65 for all concentrations.

[0301] Compression testing was performed on a second batch of FibGen (100 mg / ml + 2.5 mg / ml genipin). A representative stress-strain curve for FibGen hydrogel has three different zones of compressive behavior, defined by a different modulus (i.e., change in slope). Zone 1 is defined for a strain from 0 to 0.05 where the stress increases linearly from 0 MPa to -0.25 MPa. Zone 2 is defined for a strain from 0.05 to 0.1 where the stress increases linearly from -0.25 MPa to 0.75 MPa. Zone 3 is defined for a strain from 0.1 to 0.15 where the stress increases linearly from 0.75 MPa to 1.5 MPa. The results for the first batch are shown below in Table 13, and the results for the second batch are shown below in Table 14.Table 13Docket 88800730-000525Table 14

[0302] Additional mechanical tests were conducted for different concentrations of FibGen. Reliable and reproducible data was established, showing significant increases in compressive modulus with increasing concentration of genipen. There is no significant difference between FibGen80 and FibGen 100 at Zone 2 and 3.

[0303] Methacryl ated hyaluronic acid (HAMA) hydrogel has been identified as an good choice for filling the cartilage domain of NOVAJoint. In vivo studies with human synovial MSCs show robust cartilaginous tissue formation and an increase in chondrogenic markers compared to FibGe HA hydrogels. Both in vitro and in vivo studies with HAMA have shown that it promotes chondrogenesis from iPSC-derived chondroprogenitor (iCP) cells.

[0304] Cartilage cell seeding in hydrogel was performed. To begin, cartilage cell clusters are mixed with methacrylated hyaluronic (HAMA) acid (1%) and LAP photoinitiator (0.03%). LAP has been shown to be safe in human-use devices. The cells are mixed with the HAMA and LAP at 20 million cells / mL. The cell and hydrogel mixture is placed into custom molds (FIG. 40), and the femur and tibia inserted into the mold so that the cartilage region is submerged for 1 minute. Then, the components are removed and UV-cured for 2 minutes with a lamp (UV 405 nm), all inside a biological safety cabinet to maintain sterility. No residual LAP or HAMA remains after the UV crosslinking process. The cells penetrate into the scaffold and remain highly viable as shown by live / dead staining.

[0305] The chondrogenic differentiation process is shown below in Table 16.Docket 88800730-000525Table 16

[0306] In vivo assessment of chondrogenic stability was conducted. ASC-derived chondrogenic pellets were cultured in 384 well format for 3 weeks at 10K cells / pellet. Pellets were encapsulated in HAMA gel and implanted into SCID mice for 3 weeks. Gene expression of Sox9, AC AN and CollOal was determined. At 3 weeks gene expression results showed maintenance of chondrogenic markers but also upregulation of Col lOal . ASC pellets showed positive staining for alcian blue.

[0307] EXAMPLE 6 - Osteoinduction

[0308] The bone component of the biological knee implant is developed, demonstrating osteoinductive and osteointegrative properties to ensure bone growth while gradually degrading with time. The requirements are that this portion of the implant maintain mechanical stability through integration with the host bone tissue and allow for load-bearing. This allows for the incorporation of bioceramics and other thermally-stable factors into the 3DP fibers of the porous scaffold at the time of printing (no post-processing required). 3D printed polymer composites are fabricated consisting of polymers and biphasic bioceramics prepared from nano-hydroxyapatite (HA) and nano-P-tricalcium phosphate (0-TCP), which allows for slow dissolution of calcium and phosphate ions and apatite formation to enhance osteogenesis, osteoconduction and integration. HA and TCP have proven biocompatibility and are widely used as FDA-approved biomaterials. In addition, a novel approach to providing osteoinduction is the use of a biodegradable,Docket 88800730-000525bifunctional bioceramic component within the slow-degrading PCL, PLA, or PLA+PBAT fiber. Nano-zinc oxide (ZnO) can be embedded into the slow-degrading polymer fiber, allowing for slow-dissolution of zinc, promoting osteogenesis - stem differentiation, and providing piezoelectric activity allowing for the generation of electrical activity, which promotes bone formation, without the use of external electrodes. ZnO is considered safe / non-toxic (on the FDA's list of generally recognized as safe (GRAS) substances). The composites containing bioactive ceramics (HA / TCP) and / or ZnO provide a novel, biodegradable scaffold in combination with osteogenic cells for rapid bone formation and integration.

[0309] 3D printable polymer-ceramic composites for bone domain scaffolds were fabricated as follows. Add 15% w / v PCL to chloroform and stir for 40 minutes, add ZnO nanoparticles, and stir an additional 30 minutes. Sonicate the mixture overnight. Perform homogenous dispersion of nanoparticles. Cast in glass dish for solvent evaporation for three days. Cut dried PCL + ZnO composite into pellets for filament making with 3Dp and / or direct bioprinting.

[0310] 3D printable polymer-ceramic composites (PCL-ZnO - lOwt. %) for bone domain scaffolds were printed on two scales (FIG. 41A, FIG. 41B, FIG. 41C and FIG. 41D), both showing good print fidelity and homogeneity. Printing speed was 4 mm / second, printing pressure was 85 psi, layer height was 0.1 mm, and the printing nozzle was 30G. To further refine the process, the pellet-making and 3DP processes is tested with increased ceramic wt.%, the pellet-making and 3DP processes are tested with various polymer-ceramic mixtures, and mechanical testing is conducted on 3DP polymer-ceramic test specimens. Increasing ceramic wt.% is important to promote osteogenesis and increase scaffold strength, but high levels of ceramic may result in undesirable brittleness.

[0311] Additional bone composite scaffolds were 3D printed using a BioPrinter for characterization and in vitro studies. One of the scaffolds was PCL-10% ZnO (FIG. 42A), and another scaffold was PCL-10%ZnO+20% 20 / 80 HA / TCP (FIG. 42B). The overall dimensions were 6.5 mm x 6 mm x 1.2 mm, and the pore sizes were 200-400 pm. Bone composite scaffolds were characterized by x-ray diffraction and attenuated total reflectance-Fouri er transform infrared (ATR-FTIR) spectroscopy to ensure that there were no changes in composition / structure due to the solvent process. X-ray diffraction spectra for PCL solvent blended with ceramic nanoparticles (ZnO, ZnO + HA, ZnO + P-TCP, ZnO + HA / TCP) compared to the individual XRD spectra forDocket 88800730-000525-TCP, ZnO, HA, and PCL show that the composite materials have signals for all incorporated materials, thus indicating the solvent casting method was successful for creating polymer-ceramic composites and ensure there were no unanticipated changes in composition / structure due to the solvent casting process. ATR-FTIR spectra for PCL solvent blended with ceramic nanoparticles (ZnO, ZnO + HA, ZnO + P-TCP, ZnO + HA / TCP) compared to the individual spectra for P-TCP, ZnO, HA, and PCL show that the phosphate peaks associated with P-TCP and HA are present in the polymer composites containing those nanoparticles and ensure there were no unanticipated changes in composition / structure due to the solvent casting process.

[0312] Raw materials, solvent-casted pellets, and bioprinted specimens were characterized by ATR-FTIR that showed no unanticipated changes in composition / structure due to printing process. Measurements were taken on nanoparticles (ZnO, P-TCP and HA), PCL with wt. % nanoparticles, and PCL alone. ATR-FTIR spectra for PCL solvent blended with ceramic nanoparticles (ZnO, ZnO + HA, ZnO + P-TCP, ZnO + HA / TCP) before and after 3D printing show that the phosphate peaks associated with P-TCP and HA are present in the polymer composites containing those nanoparticles after printing, thus showing successful incorporation of the ceramic nanoparticles into the printed scaffolds and ensure there were no unanticipated changes in composition / structure due to the printing process. XRD spectra for PCL solvent blended with ceramic nanoparticles (ZnO, ZnO + HA, ZnO + p-TCP, ZnO + HA / TCP) before and after 3D printing show that the crystalline peaks of ZnO, HA, and P-TCP are present in the polymer composite scaffolds, thus showing successful incorporation of the ceramic nanoparticles into the printed scaffolds and ensure there were no unanticipated changes in composition / structure due to the printing process.

[0313] The porosity, strut size and pore size were determined for a variety of bone composite scaffolds using polymers and polymer-ceramic composites. The polymers tested were PCL and solvent-casted PCL, and the polymer-ceramic blends tested were PCL- 10% ZnO, PCL-10% ZnO-20% TCP, PCL- 10% ZnO-20% HA, PCL- 10% ZnO-20% HA / TCP (20 / 80), PCL-30% TCP, PCL-20% HA, and PCL-20% HA / TCP (20 / 80). The porosity results, the strut size results, and the pore size results show consistent printing results across various polymer-ceramic composite compositions.

[0314] Energy Dispersive X-ray Analysis (EDXA) mapping further confirmed distribution of ceramic nanomaterials (PCL- 10% ZnO-20% P-TCP) on the scaffold surface. SurfaceDocket 88800730-000525characterization of 3D printed PCL-ZnO (10 wt.%)- 0-TCP (20 wt.%) scaffolds with scanning electron microscopy (SEM) and energy-dispersive x-ray spectroscopy (EDX) show that carbon, oxygen, phosphorous, calcium, and zinc are dispersed evenly across the scaffold surface, indicating a homogeneous incorporation of the ceramic nanoparticles into the polymer. The samples were carbon (C) coated - C wt. % is not from composite alone.

[0315] Mechanical testing on the compressive properties of bioprinted PCL and bioprinted PCL + TCP + ZnO (70 / 20 / 10 wt. %) polymer-ceramic composites was performed. All samples were printed with 70% infdl and tested at high strain rate (80% / s). The bioprinter software does not accurately reproduce the targeted infill density (z.e., porosity), producing significantly poorer mechanical properties in compression than anticipated for actual 70% infill scaffolds. The actual infill for both samples was -40%. Tensile testing of 70% infill 3D printed PCL and PCL-ZnO-TCP composite scaffolds showed that PCL-ceramic composites have large variability in their mechanical properties, so no statistically significant improvement was observed. The yield stress, Young’s modulus, and Toughness for both materials along with their standard deviations are in Table 17 below. The results show a very large variability in properties in the PCL-ceramic composites, arising from the bioprinting method where different size and weights of scaffolds were used in these tests. No statistically significant improvement was observed.Table 17

[0316] Thermogravimetric analysis (TGA) of the 3D printed bone composite scaffolds confirms the weight-percent of nanoparticles in the composite after processing is consistent with the initial mixed composition.

[0317] Cytotoxicity of bone composite scaffolds was investigated. Live / dead staining and adhesion of cells to PCL and composite (PCL + 10% ZnO + 20% 0-TCP) scaffolds. Scaffolds were preconditioned in media containing serum overnight, and the cells were cultured for two days. Live-dead staining (live = green, dead = red) showed viable cells and attachment.Docket 88800730-000525

[0318] PLA samples printed with the 3D printer and bioprinter have significantly different properties. Bioprinter software does not accurately reproduce the infill percentage from the Prusa-Slicer CAD models. For 70% infill CAD models, the 3D printer produced an actual infill of -67%, whereas the bioprinter produced an actual infill of -40 %. PCL from the bioprinter was much weaker that PCL from the Prusa printer, likely due to noted structural differences. All samples were printed with 70% infill and tested at high strain rate (80% / s). The compressive properties for PCL with (nominal) 70% infill at high strain rate (80 / s) on the Prusa printer resulted in a yield stress of 9.6 + 0.1 MPa, Young’s modulus of 120 + 3 MPa, and toughness of 11.3 + 0.4 MPa. The compressive properties for PCL with (nominal) 70% infill at high strain rate (80 / s) on the bioprinter resulted in a yield stress of 2.3 + 0.1 MPa, Young’s modulus of 25 + 2 MPa, and toughness of 5.7 + 0.6 MPa.

[0319] 3D printed polycaprolactone (PCL) composites were fabricated and characterized. PCL was combined with a range of bioceramics including zinc oxide (ZnO), hydroxyapatite (HA), beta-tricalcium phosphate (0-TCP) and 20 / 80 (20 wt.% / 80 wt.%) HA / 0-TCP. The bioceramics were nanoparticles (-100 nm) to allow for more rapid dissolution / degradation while embedded in a slow-degrading polymer. HA and P-TCP were incorporated to promote the formation of apatite, which is important for osteointegration. All bioceramics have been shown to affect VEGF secretion by stem cells and other cell types, which is important for neovascularization. The 20 / 80 HA p-TCP ratio was identified, both in vitro and in vivo, to promote stem cell osteogenic differentiation and bone tissue formation. Surface analysis techniques, x-ray diffraction (XRD), attenuated reflectance - Fourier transform infrared spectroscopy (ATR-FTIR) and thermogravimetric analysis (TGA) demonstrated the presence of the nanoceramics - surface and bulk and loading efficiency (>90%). The average pore size was approximately 250 pm and the strut size was 350 pm.

[0320] Print fidelity is consistent between PCL and PCL-ceramic bioprints. Although the bioprinter specimen infills are inconsistent with the scaffold CAD models and 3DP specimens, there is consistency among bioprinted polymer and polymer-ceramic composites. Optical images and SEM images of PCL (FIG. 43, top) and PCL-ceramic (FIG. 43, bottom) specimens manufactured with the bioprinter. Both prints show good print fidelity and minimal defects. The increase in visible surface roughness of the PCL-ceramic composite is expected due to the ceramicDocket 88800730-000525nanoparticles. Statistical analysis of the strut sizes and pore sizes of the PCL and PCL-ceramic prints shows no significant change in sizes.

[0321] Bone domain composite materials were investigated. A composite of Ecovio T2308 + 30 wt.% hydroxyapatite / p-tricalcium phosphate (20 / 80) [T2308+HT] was identified as a good choice for NOVAJoint. Results from tensile and compressive testing of 70% infill alternating triangular specimens indicate the material has sufficient properties to sustain mechanical loading. Results from cellular studies show that the T2308+HT scaffolds promote cell adhesion, growth, and osteoinduction. Studies with 40 wt.% HT composites are being conducted to see if they further improve osteoinduction while maintaining mechanical properties.

[0322] Bone cell seeding was carried out as follows. First, the sterilized scaffolds are plasma-treated before cell seeding. Next, fibronectin is adsorbed onto the bone region of the implant by placing each implant (femur and tibia) and each test coupon into a custom chamber, filling it with the fibronectin solution (100 pg / mL in PBS) until the bone portion is covered, and incubating for 1 hour. After incubation, the implant is air-dried in a biological safety cabinet for 3 hours. Next, bone cells (10M cells / mL) are suspended in MOPS media (500 ml DMEM High glucose; 0.1 pM; Dexamethasone; 50 ng / ml ascorbic acid; IX ITS; 2 mM L-glutamine; 1 mM sodium pyruvate; 100 U / ml; IX Non-essential Amino Acids) and loaded into a 10 m leur-lock syringe. The coated implants and test coupons are each reinserted into their custom chambers and loaded with the cell suspension from the bottom loading port of the chamber until the bone region is covered. Finally, the chambers are placed on an orbital shaker for 3 hours at 120 rpm at 37°C in a 5% CO2 incubator. Rendering and photo of chamber prototype to seed bone cells onto tibial component are shown in FIG44A, FIG. 44B and FIG. 45C. Rendering and photo of chamber used to seed bone cells onto femoral component are shown in FIG. 45 A, FIG. 45B, FIG. 45C and FIG. 45D. Cell seeding near interface of tibial region was performed. Bone cell seeding in cross-section of tibial keel was performed.

[0323] Bone domain scaffold cell seeding methods have good loading efficiency. Dynamic cell loading conditions: gentle agitation using an orbital shaker (3 hours); cell loading media volume: 100 pL; cell loading densities: 5 million / ml and 10 million / ml. Increased cell loading with higher cell seeding density observed, with minimal loss of cells.

[0324] Mechanical properties of acellular bone and cartilage scaffolds, before and after sterilization was investigated. Ecovio T2308 compressive properties were evaluated on cylindricalDocket 88800730-000525samples and tensile testing on dog bone samples pre- and post-sterilization using an Tnstron mechanical test system. The measured yield stresses for both compression and tension are below the peak tensile and compressive stresses as determined through finite element (FE) models of NOVAKnee-T designs. No change in tensile or compressive mechanical properties of Ecovio T2308 after ethylene oxide sterilization was noted for the cartilage scaffold.

[0325] Osteochondral implant mechanical evaluation in loading bioreactor was conducted. Cell seeded human (chondroprogenitors and osteoprogenitors) cylindrical osteochondral constructs were cultured for up to 28 days with and without applied shear loading (1 mm / s reciprocal sliding against 14 inch titanium ball bearing, 10 MPa load for 3 hours / day). Mechanical testing showed that mechanical properties were maintained with and without loading at multiple time points in culture. As anticipated, mechanical properties were unchanged with culture, as the stiffness of constructs is dictated by the scaffold which is orders of magnitude stiffer than HAMA-cells (and any ECM elaborated) in the pores. Cells and extracellular matrix found through osteochondral scaffold.

[0326] The osteogenic differentiation process is shown below is Table 18.Table 18

[0327] In vivo assessment of osteogenic stability was investigated. ASCs were expanded for 1 week, cultured in osteogenic media for 72 hours and then seeded on T2308+HT scaffolds at a density of 10 M cells / ml. Cell-seeded T208+HT scaffolds were implanted in SCID mice subcutaneously (4 constructs per mouse) and harvested after 3 weeks. Gross morphology andDocket 88800730-000525body weight measurements indicate there were no deleterious effects. Alizarin red staining increased in the ASC-seeded osteogenic scaffolds after 3 weeks in vivo. ASCs from four patients were differentiated for 3 weeks in osteogenic media followed by seeding in scaffolds and implantation in SCID mice. After 3 weeks of implantation, two ASC samples maintained osteogenic gene expression while two showed equivalent or reduced expression to naive ASCs. Reduced in vitro culture time (less than 3 weeks) may improve osteogenic outcomes after in vivo implantation.

[0328] EXAMPLE 7 - Vascularization

[0329] With regard to angiogenesis, the bony portion of the presently disclosed biologic living TKA joint is 3D printed containing PCL or other biopolymers of interest with bioactive ceramics (HA, TCP, ratio of HA / TCP and / or ZnO). With this approach, strategies for angiogenesis include: A) Porous biomaterials consisting of HA and TCP that have been shown to upregulate vascular endothelial growth factor (VEGF) expression in vitro and in vivo and promote neovascularization in both ectopic and bone defect models in vivo, B) 10 wt.% of ZnO in a ZnO— PCL composite can promote VEGF expression in human mesenchymal stem cells (MSCs); C) The differentiated cells, osteoblasts (OBs) and hypertrophic chondrocytes (HCs), that are seeded onto the implants are known to secrete pro-angiogenic factors, such as VEGF, which is required for blood vessel invasion and subsequent bone formation. Changes in seeding density and stem cell -derived endothelial cells can also be used in combination with the undifferentiated and differentiated cells (ASC, HC, and OB) to promote vascularization.

[0330] In certain embodiments, the scaffold can be designed to include growth factors or one or more antibacterials or other medicinal agents or materials. These agents or materials may be incorporated directly into the scaffold, into the material injected into the scaffold, or included in microspheres that are built into or injected into the scaffold. In certain embodiments, the microspheres can be designed with a specific degradation timeline, or multiple degradation timelines for different portions or subgroups of the microspheres. For example, all of the microspheres may release the agent(s) or material(s) in 1 month, in 2 months, in 3 months, in 4 months, in 5 months, in 6 months, in 7 months, in 8 months, in 9 months, in 10 months, in 11 months, in 12 months, in 13 months, in 14 months, in 15 months, in 16 months, in 17 months, or in 18 months, from the date of implantation. In certain embodiments, different portions of groups of microspheres may be released sequentially or periodically, such as a first portion released in 1Docket 88800730-000525month, in 2 months, in 3 months, in 4 months, in 5 months, in 6 months, in 7 months, in 8 months, in 9 months, in 10 months, in 11 months, in 12 months, in 13 months, in 14 months, in 15 months, in 16 months, in 17 months, or in 18 months, and a second portion or subgroup released 1 month after the first portion, 2 months after the first portion, 3 months after the first portion, 4 months after the first portion, 5 months after the first portion, 6 months after the first portion, 7 months after the first portion, 8 months after the first portion, 9 months after the first portion, 10 months after the first portion, 11 months after the first portion, or 12 months after the first portion. There can be other portions or subgroups of the microspheres that are released after the second portion, and so on, including in a sequential manner.

[0331] In vitro vascularization evaluation of bone domain composites was investigated. Human umbilical vein endothelial cells (HUVECs) were cultured on Ecovio T2308 + 30 wt.% HA / TCP (20 / 80) [T2308+HT] bone domain scaffolds for 14 days as an in vitro assessment of angiogenesis. After two weeks, an increase in cell numbers was observed on both T2308 and T2308+HT scaffolds and samples stained positive for VEGF-A and CD31. An increased production of VEGF-A from ASC-derived osteoblasts (ASC-OBs) on T2308+HT scaffolds was observed compared to T2308 alone. VEGF is a known chemoattractant of endothelial cells and promotes vascular invasion, therefore suggesting that, when the endothelial cells get there via migration, the phenotype and function will be maintained for endothelial cells on the bony scaffold and may enhance endothelial proliferation and angiogenesis once implanted in vivo. Actin, DAPI, and VEGF-A staining for HUVECs cultured on T2308 and T2308+HT scaffolds for 14 days was examined. Actin, DAPI, and CD31 staining for HUVECs cultured on T2308 and T2308+HT scaffolds for 14 days was examined. Actin, DAPI, and VEGF-A staining was examined. Sustained VEGF and CD31 expression for HUVECs (Day 14) on T2308 and T2308-HT30 scaffolds was observed.

[0332] Example 8 - Nonclinical Testing

[0333] Overview

[0334] The nonclinical testing program is designed to consider both the device and biological components. In vitro and small animal testing will be conducted to evaluate the properties of the presently disclosed biologic living TKA joint (NOVAIoint) by region and on multilayer osteochondral test coupons. Degradation and mechanical testing will also beDocket 88800730-000525performed. Jn vivo studies will evaluate construct maturation and safety. Both a toxicology study in a rodent model and a large animal OA model in canines are proposed.

[0335] The Design Control process is being utilized as part of the development efforts. Potential hazards from NOVAJoint will be identified and studies will be conducted to understand the risk of these hazards, potentially redesigning NOVAJoint if results are determined to be unacceptable. Degradation assessment, cytotoxicity, particulate formation, and implantation testing are planned. Pyrogenicity will be routinely tested as part of lot release. Tests related to sensitization, genotoxicity and carcinogenicity will be performed, per ISO10993.

[0336] The target properties are such that the peak contact stresses remain below the yield stress of the scaffold material and that the peak tensile stresses remain below the ultimate stress of the scaffold material in the final scaffold designs. The inventors target NOVAKnee-T to allow for improved knee functionality for daily activities approximately at 6 weeks post-op. Detailed methodologies will be provided in the future. Also, the inventors will provide a rationale for the how the assessments address the structural function of cartilage and bone. Both in vitro and in vivo studies will be performed to evaluate the proposed designs against the target properties. The target properties are such that the peak contact stresses remain below the yield stress of the scaffold material and that the peak tensile stresses remain below the ultimate stress of the scaffold material in the final scaffold designs. NOVAKnee-T is also targeted to allow for improved knee functionality for daily activities approximately at 6 weeks post-op. Both in vitro and in vivo studies will be performed to evaluate the proposed designs against the target properties.

[0337] Proof-of-Concept Studies Completed to Date

[0338] Early proof-of-concept (POC) studies have been executed for NOVAJoint that are focused on scaffold design and cell manufacturing. The dimensions and curvature of the femoral and tibial components are defined (FIG. 46A, FIG. 46B, FIG. 47), along with delineation between cartilage and bone regions. An alternating triangular infill pattern has been selected (FIG. 49). The feasibility of including a sacrificial polymer scaffold has been shown to enable three-dimensional printing of NOVAJoint. The exact printing pattern and polymer compositions are being optimized, as it impacts mechanical properties (FIG. 47 and FIG. 48). It was learned that plasma treatment may be needed to enable cell seeding. Water contact angle testing for plasma-treated Ecovio scaffolds shows increased wettability of the scaffold with increased time of plasma treatment. With one minute of treatment, a water droplet remains completely on the surface of theDocket 88800730-000525scaffold with a high contact angle. With three minutes of treatment, a water droplet remains on the surface of the scaffold but flattened to a lower contact angle. At five minutes of treatment, a drop of water flows completely through the porous scaffold. The process of isolating ASCs from lipoaspirate has been performed a number of times, and the resulting cells have shown a consistent output. Optimization of printing hydrogel onto the scaffold is ongoing. Preliminary degradation and cytotoxicity studies are ongoing, although data is still pending.

[0339] Custom filaments have been used to print various patterns, with methods being developed to improve consistency of materials and printing. FTIR spectra of Ecovio T2308 compared to the independent spectra of PLA and PBAT was applied and a novel algorithm was developed to determine and / or validate the relative weight of each polymer component in the blend. The algorithm normalizes the FTIR spectrum by the area under the curve, and then applies a best fit algorithm of the polymer blend spectra against the sum of one or more spectra of the defined subcomponents. Results from this method indicate that the tested sample of Ecovio T2308 has a composition of 85 wt.% PLA and 15 wt.% PBAT. One iteration of the FTIR analysis algorithm included an unknown entity to provide information on whether reactive components or unknown contaminants may be present in the polymer blend. Reactive components were found to be below level of detection and no major contaminants were identified. Measurements of infill density via mass / volume measurement are compared to measurements obtained via microCT across multiple 3D printed samples as a validation of print fidelity and use of microCT for infill density analysis after printing. The outcomes of these various measurements and calculations are in Table 19 and Table 20 below. Prusa “Ideal Conditions” calculations are based on theoretical mass (0.86 g / sample for lx multiplier and theoretical volume (pi*(14.38 / 2)2 * 6) = 973 mm3. Prusa “Real Conditions” calculations are based on theoretical mass (1.05 g / sample for 1.23x multiplier for triangular pattern) and theoretical volume (pi*(14.38 / 2)2 * 6) = 973 mm3. The microCT calculations are performed using a threshold of 14 / 1000.Table 19Docket 88800730-000525Table 20

[0340] PLA, Ecovio T2308, and T2308 + HA (30 wt.%) scaffolds were 3D printed using various values for the Prusa extrusion multiplier setting. The multiplier is intended to compensate for differences in melt viscosity among different polymer materials. For all three materials, there is a linear increase in actual infill density as a function of the multiplier. To determine the appropriate multiplier for each material, the linear relationship is used to determine where the measured infill will reach 70% (desired nominal infill). The multiplier determined to be necessary for 70% infill density for each material is l.lOx for PLA, 1.23x for T2308, and l.Ox for T2308+HA.

[0341] To evaluate print fidelity, microCT was used to measure fiber widths throughout various sections of a 3D printed T2308 cylindrical scaffold. The scaffold was printed with a 70% alternating triangular (30° / 60°) infill pattern and 1.23x multiplier. The nominal width, average width, and standard deviation of width for the top, middle, and bottom of the sample, as well as the overall sample are given in Table 21 below. There is a significant decrease in fiber width for the top layer of the sample compared to the middle and bottom, with no significant difference between the middle and bottom nor the top, middle, and bottom relative to the overall scaffold.Docket 88800730-000525Table 21

[0342] A study was conducted to evaluate the impact on extruder cleaning via “cold pull” on eliminated defects within a print sample. MicroCT analysis of a sample printed before cold pull showed microscopic defects in 15.4% of layers in a porous scaffold. Nozzle residue is visible on the filament used for a first pull, and a repeated cold pull shows residue is reduced on the second pull. Once the cold pull process has been repeated to sufficiently remove residues, another sample is printed. In two samples printed after cold pull, the first had no defects and a second has only one defect, which is calculated to only 1.51% of layers containing defects.

[0343] With regards to using iPSCs, the inventors have successfully grown the cells and differentiated them, showing low residual iPSCs after culture. Bone cells have been seeded onto scaffolds with long-term viability and functionality shown. Chondrogenic cells in hydrogel have been bioprinted with success. The mechanical properties of the hydrogel with and without GAG mimetics have been evaluated. The compressive moduli of FibGenlOO and FibGen80 with genipin concentrations of 2.5 mg / mL, with or without 1% GAG mimetics (partially sulfated dextran (Dex) or fully sulfated cellulose (NaCS)) were measured for three different zones of compressive deformation. It was found that the addition of GAG mimetics does not significantly alter the mechanical properties of the hydrogel at either concentration, with moduli of ~5 MPa in Zone 1 for both FibGen concentrations, ~15 MPa in Zone 2 for both FibGen concentrations, and ~35 MPa for FibGenlOO and ~25 MPa for FibGen80 in Zone 3. The infill of hydrogel has been evaluated without cells after printing onto the scaffolds. Micrographs of 3-minute 02 plasma treated PLA / PBAT scaffold showed homogenous coating of FibGen and its infiltration into the scaffold. Infiltration depths measured at multiple locations of 4 different scaffolds show consistent measures between 2-2.5 mm.

[0344] Additional proof-of-concept studies have been performed. The inventors have performed preliminary degradation and cytotoxicity studies. The inventors have performed mechanical tensile testing and compressive testing of proposed alternating triangular infill patternDocket 88800730-000525scaffold morphology 3D printed with Ecovio T2308; and have conducted compressive testing on polymer-ceramic composites with the same scaffold morphology. Tensile testing was performed on porous scaffolds of T2308 printed with either alternating (30 60°) triangular pattern or alternating (07457135790°) rectilinear pattern. Both patterns were printed with a 70% infill density. The toughness for both infill patterns is approximately 0.175 MPa while the ultimate tensile strength is significantly increased in the triangular pattern at ~15 MPa compared to 9 MPa for the rectilinear pattern scaffold. Compressive testing was performed on porous scaffolds of T2308 printed with either alternating (30760°) triangular pattern or alternating (07457135790°) rectilinear pattern. Both patterns were printed with a 70% infill density. The average yield strength was ~33 MPa for the triangular pattern scaffold and ~27 MPa for the rectilinear pattern scaffold. The average toughness was ~48 MPa for the triangular pattern scaffold and ~40 MPa for the rectilinear pattern scaffold. Compressive testing was performed on porous scaffolds of T2308 with either 10 wt.% or 30 wt.% hydroxyapatite (HA). For the T2308 + 30% HA material, testing was performed on both alternating triangular and alternating rectangular infill pattern scaffolds. All scaffolds were printed with a 70% infill density. The average measured yield strength and toughness for each scaffold is in Table 22 below. An increase in HA decreased the yield strength in rectilinear scaffolds, but the yield strength was recovered when the scaffold was printed with a triangular pattern. An increase in HA increased the toughness in rectilinear scaffolds, with the toughness staying approximately equal across rectilinear and triangular scaffolds for 30% HA. (FIG. 108).Table 22

[0345] The inventors also performed preliminary friction testing of the hydrogel. Representative creep displacement curve for a friction test of FibGen coated on a T2308 scaffold starts at a creep displacement of 0 pm which steeply drops to — 2200 pm after 20 cycles then holds steady up to 110 cycles. A representative friction coefficient curve for a friction test ofDocket 88800730-000525FibGen coated on a T2308 scaffold starts at friction coefficient of 0.095 which steeply drops to approximately 0.055 and remains at that level with small fluctuations (<.001 ) up to 100 cycles. . The inventors also performed POC mechanical testing of the first 3D printed anatomical NOVAJoint construct, which showed no sudden drop in load-displacement.

[0346] Most recently, the inventors implanted a 3D printed NOVAJoint into a human cadaver (with advanced OA) using the Zimmer Persona System and Smartnails, and performed mechanical testing at full extension, 30°, 60° and 90° with a force of 90 N then 700 N (FIG. 50A, FIG. 50B, FIG. 50C, FIG. 50D, FIG. 51 A, FIG. 5 IB, FIG. 51C and FIG. 5 ID). No failure, obvious damage, or implant shifts were observed during motion, and the patella tracked well.

[0347] Species justification

[0348] Most of the animal studies proposed in this program utilize the athymic rat. The subcutaneous athymic rat model (male) is chosen as it is a well-established model fortesting tissue engineered scaffolds. The rat strain is ideally suited to studies with cross-species implantation of tissue engineered scaffolds. Utilization of this model will avoid an immune response to the implanted scaffolds and cells. Minimal immune response is achieved due to the rats’ lack of T cells. Still, athymic rats produce macrophages which produce a cytokine-rich environment that cannot be achieved in vitro. This is the smallest animal model in which the response can be evaluated because the rat dorsum is of the adequate size for the proposed studies. Also, rats can accommodate multiple scaffolds subcutaneously within a single animal, which in turn minimizes the number of animals required.

[0349] Wild-type immunocompetent rats will be used to evaluate immunogenicity of cell free NOVAJoint scaffolds in a subcutaneous implantation mode. 3D printed scaffolds will be subcutaneously implanted and evaluated upon resection at defined timepoints for histological analysis to evaluate potential of a sustained foreign body or immune response. This model is well established as a method to evaluate immunogenicity of novel materials.

[0350] Additionally, one study is proposed in the (Severe Combined Immunodeficiency) SCID mouse to evaluate osteoinduction. The SCID mouse model is used as it is a well-established model for evaluating biomaterials, stem-cell induced bone formation and as supportive scaffolds for bone ingrowth. The SCID mouse is an immunocompromised mouse that is used for crossspecies implantation of cells for the evaluations of their potency. The tissue engineered scaffolds can be evaluated ectopically (subcutaneously) for bone induction in this mouse model using up toDocket 88800730-0005256 implants per mouse as the back of the mouse has a large surface area for implantation of scaffolds to adequately assess both qualitatively and quantitatively bone formation and neovascularization in the implants.

[0351]

[0352] Additional proof of concept studies have been completed. Completed in vitro mechanical and material testing for NOVAJoint are shown in Table 23.Table 23

[0353] Completed in vitro biological studies for NOVAJoint (NJ), autologous NOVAKnee (NKT), and allogeneic NOVAKnee (NKL) are shown in Table 24.Table 24Docket 88800730-000525

[0354] Completed in vivo biological studies for NOVAJoint, NOVAKnee-T, and NOVAKnee-L are shown in Table 25.Table 25

[0355] Completed testing on anatomically shaped NOVAKnee scaffolds is shown in Table 26.Docket 88800730-000525Table 26

[0356] Proposed In Vitro and In Vivo Testing

[0357] A series of in vitro and in vivo studies will be conducted. Proposed in vitro testing studies for NOVAJoint, NOVAKnee-T, and NOVAKnee-L are shown in Table 27.Docket 88800730-000525Table 27Docket 88800730-000525

[0358] Proposed in vivo studies for NOVAJoint, NOVAKnee-T, and NOVAKnee-L are shown in Table 28.Table 28Docket 88800730-000525Docket 88800730-000525Docket 88800730-000525Docket 88800730-000525

[0359] Additional Proposed Large Animal Studies

[0360] An additional large animal study is proposed to ensure the surgical technique is suitable. The inventors believe that a canine osteoarthritis model is the best approach to evaluating NOVAJoint. The canine meniscus release (MR) model provides the optimal preclinical animal model for total joint replacement for knee osteoarthritis (OA) based on its anatomical, biological, biomechanical, and clinical relevance. Other than non-human primates, the canine stifle (knee) most closely resembles the human knee anatomy, biology / physiology, and functional biomechanics. This is clearly manifested in the etiopathogenesis of knee OA in clinical canine patients, which occurs as the result of cranial (anterior) cruciate ligament tears, meniscus deficiency, overuse, trauma, and aging.

[0361] The canine MR model consistently induces symptomatic whole-joint OA, including full-thickness cartilage loss, synovitis, effusion, osteophytosis, and subchondral bone changes that mimic those diagnosed in human gonarthrosis. Diagnostic imaging (radiography, CT, MRI, arthroscopy) is readily performed in dogs and supports each of these pathologic findings. Consistent symptoms in MR dogs include pain, lameness, and decreased range of motion, which can be accurately assessed using validated outcome measures. Further, this model allows for use of clinically relevant pre- and post-operative management strategies including “prehab”, orthotics, analgesics, anti-inflammatories, orthobiologics, nutritional alterations, activity modifications, and full-spectrum physical therapy.

[0362] Importantly, dogs are the only animal species with a commercially available total knee arthroplasty system and in which biologic joint resurfacing procedures have been performed, both with outcomes in clinical canine patients documented. As such, this model provides the onlyDocket 88800730-000525validated method for clinical applicable comparisons to the standard-of-care comparators for NOVAJoint.

[0363] To evaluate NOVAJoint in an anatomically relevant size, canine-sized NOVAJoint constructs will be generated and tested in vitro for mechanical strength. Then, cadaveric study will be conducted to ensure the surgical technique is suitable. After demonstrating consistency of the canine OA model, a pivotal large animal study will be performed in 12 OA dogs as described in Table 29. The cells to be used in this study will be generated using the process to be transferred to our manufacturing site. Any changes that occur will be discussed to evaluate...

Claims

Docket 88800730-000525CLAIMS1. A biologic living joint replacement device, comprising:a first component comprising a first bone layer and a first cartilage layer; anda second component comprising a second bone layer and a second cartilage layer, wherein the first bone layer and the second bone layer each comprises a porous structure made of biocompatible and bioresorbable polymer, polymer blend, or polymer-ceramic composite and stem cell-derived bone cells, and wherein the first cartilage layer and the second cartilage layer each comprises a porous structure made of biocompatible and bioresorbable polymer or polymer blend and stem cell-derived cartilage cells.

2. The device of claim 1, wherein the device is produced by 3D printing technology.

3. The device of claim 1, wherein the shape of the device is produced by scanning existing three part implants, and combining the three parts into two components comprising the first component and the second component.

4. The device of claim 1, wherein the device shape is modified using in silico finite element modeling to reduce stresses and improve surface conformity.

5. The device of claim 1, wherein each of the first component and the second component comprises a cartilage domain, a bone domain, and an impermeable layer, and wherein the cartilage domain and the bone domain are separated by the impermeable layer.

6. The device of claim 1, wherein the first component further comprises an articular surface, wherein the second component further comprises an articular surface, wherein a cartilage subdomain is added to an articular surface of each of the first component and the second component to control lubrication properties.

7. The device of claim 1, wherein the mechanical properties of the device are designed to meet the mechanical demands of the joint under activities of daily living.

8. The device of claim 1, wherein the cartilage domain and the bone domain each include polymer blends, biomaterials, and biological factors to promote cartilage formation in the cartilage domain and bone formation in the bone domain and to promote phenotype maintenance after implantation.

9. The device of claim 1, wherein the stem cells are autologous stem cells.

10. The device of claim 9, wherein the stem cells are autologous adipose-derived stem cells.

11. The device of claim 1, wherein the stem cells are allogeneic stem cells.Docket 88800730-00052512. The device of claim 11, wherein the stem cells are multipotent allogenic stem cells.

13. The device of claim 12, wherein the stem cells are human induced pluripotent stem cells.

14. The device of claim 1, wherein the first cartilage layer and the second cartilage layer each comprises genipin crosslinked fibrin hydrogel.

15. The device of claim 1, wherein the first cartilage layer and the second cartilage layer each comprises photocrosslinkable methacrylated hyaluronic acid (HAMA) hydrogel or a genipin-crosslinked fibrin (FibGen) hydrogel.

16. The device of claim 1, wherein the first bone layer and the second bone layer each comprises a polymer-ceramic composite.

17. The device of claim 1, wherein at least one of the first cartilage layer or the second cartilage layer comprises a genipin-crosslinked fibrin (FibGen) hydrogel or a photocrosslinkable methacrylated hyaluronic acid (HAMA) hydrogel configured to delay degradation and provide improved compressive mechanical properties suitable for cartilage formation in vivo.

18. The device of claim 1, wherein the first bone layer and the second bone layer each comprises a polymer-ceramic composite selected from polycaprolactone (PCL), polylactic acid (PLA), PLA / PBAT blends, or combinations thereof with ceramic particulates including hydroxyapatite (HA), P-tricalcium phosphate (0-TCP), zinc oxide (ZnO), or combinations thereof.

19. The device of claim 1, wherein the first cartilage layer and the second cartilage layer each comprises a polymer blend comprising PLA and PBAT configured to provide increased ductility and toughness relative to PLA alone.

20. The device of claim 1, wherein at least one of the first component or the second component comprises a 3D-printed porous scaffold having an infill of about 70% to provide a porosity of about 30% with alternating raster orientations to balance pore size and strength.

21. The device of claim 1 , wherein the scaffold pore architecture comprises an alternating infill pattern or a gyroid pattern, each configured to provide comparable equilibrium compressive modulus in the cartilage or bone domain.

22. The device of claim 1, wherein the first articular surface and the second articular surface are each configured to provide increased congruence and reduced peak contact pressure relative to unmodified commercial geometries, as determined by finite element modeling of plastically deformed reference geometries.Docket 88800730-00052523. The device of claim 1, wherein the cartilage domain and the bone domain of at least one the first component or the second component are mechanically and spatially distinct from one another in a multi-layered construct configured to withstand peak loads representative of about ten times body weight without failure.

24. The device of claim 1, wherein the stem cell derived bone cells and cartilage cells are obtained from autologous adipose-derived stem cells (ASCs) or from allogeneic human induced pluripotent stem cells (hiPSCs).

25. The device of claim 1, wherein the cartilage layer comprises heparin-conjugated FibGen hydrogel or HAMA hydrogel configured to promote lubricin adsorption for boundary lubrication and reduced friction.

26. The device of claim 18, wherein the bone layer includes ZnO nanoparticles embedded within a slow-degrading polymer fiber to provide slow zinc ion release and piezoelectric activity to promote osteogenesis.

27. The device of claim 1, wherein each of the first component and the second component comprises a 3D printed scaffold, each exhibiting pore sizes of about 200-400 pm and strut sizes of about 350-450 pm to support cellular infiltration and tissue integration.

28. The device of claim 5, wherein the cartilage domain employs a ceramic-free polymer blend that is more compliant and ductile, and the bone domain employs a polymer-ceramic composite to assist in osteoinduction.

29. The device of claim 14, wherein the cartilage layer hydrogel is deposited using a dualchannel nozzle enabling in situ mixing and deposition onto the scaffold surface with controlled thickness and infiltration depth by tuning feed rate and nozzle height.

30. The device of claim 5, wherein the cartilage domain and the bone domain are selected to meet domain-specific mechanical targets derived from finite element modeling, including bonedomain tensile strength exceeding about 27 MPa and compressive strength exceeding about 63 MPa, and cartilage-domain tensile strength exceeding about 6 MPa and compressive strength exceeding about 48 MPa.

31. The device of claim 1, further comprising a multi-material scaffold, wherein the scaffold is additively manufactured using a multi-material printhead to produce a discrete cartilage domain and a discrete bone domain and an osteochondral architecture with non-planar articular contours.Docket 88800730-00052532. The device of claim 1, wherein the cartilage layer and the bone layer are configured to interact without menisci.

33. The device of claim 31, wherein the scaffold is designed to be entirely bioresorbable and non-immunogenic and chondro-inductive and osteo-inductive in vivo.

34. The device of claim 1, wherein the polymer-ceramic composite further comprises a biphasic ceramic of hydroxyapatite and P-tricalcium phosphate to promote osteoconduction and osteointegration.

35. The device of claim 1, further comprising an articular layer geometry on at least one of the first component or the second component, wherein the articular layer geometry is defined using finite element deformation to target a desired average cartilage thickness and congruent contact under physiologic load.

36. A method of preparing a biologic living joint replacement device, comprising the steps of fabricating a first component comprising a first bone layer and a first cartilage layer; and fabricating a second component comprising a second bone layer and a second cartilage layer.

37. The method of claim 36, wherein at least one of the first component or the second component is fabricated using 3D Printing.

38. The method of claim 36, wherein the first bone layer and the second bone layer each comprises a biocompatible and bioresorbable polymer-ceramic composite and stem cell derived bone cells and the first cartilage layer and the second cartilage layer each comprises a biocompatible and bioresorbable polymer blend and stem cell derived cartilage cells.

39. The method of claim 38, wherein the polymer blend is a hydrogel and the cartilage cells and hydrogel are bioprinted into cartilage layers or deposited in the layer using a dip coating method via a negative mold.

40. The method of claim 38, further comprising seeding the bone cells into the bone domain using a seeding device comprising a container configured to drive a flow of bone cells in media only into the bone regions, and preventing bone cells from reaching the cartilage region via an impermeable form fitting barrier.

41. A biologic joint replacement device, comprising:at least one of a first component comprising a first bone layer and a first cartilage layer, or a second component comprising a second bone layer and a second cartilage layer, wherein the at least one of the first bone layer or the second bone layer comprises a biocompatible andDocket 88800730-000525bioresorbable polymer-ceramic composite and stem cell derived bone cells and the at least one of the first cartilage layer or the second cartilage layer comprises a biocompatible and bioresorbable polymer, polymer blend, or polymer composite and stem cell derived cartilage cells.

42. The device of claim 41, wherein the device is secured to a bone using ancillary fixation bioresorbable pins, screws, or related fixation devices.

43. The device of claim 1, wherein the device comprises a scaffold having four layers including: a coating layer to minimize friction; a cartilage-seeding layer; an impermeable interface layer configured to reduce or prevent cell migration between domains; and a bone-seeding layer comprising a polymer-ceramic composite.

44. The device of claim 43, wherein the coating layer and cartilage-seeding layer each comprise Ecovio T2308 with about 70% infill, and the bone-seeding layer comprises Ecovio T2308 with hydroxyapatite and P-tricalcium phosphate.

45. The device of claim 1, wherein the cartilage domain and the bone domain of the scaffold is optimized to a porosity of about 30%, with an infill density of about 70%, using sequential layers with rectilinear fibers alternating by about 45 degrees to enhance resistance to tensile and compressive failure.

46. The device of claim 31, wherein the cartilage domain comprises a triangular infill pattern to increase dynamic compressive modulus while maintaining interconnectivity for cellular communication.

47. The device of claim 4, wherein finite element analysis of a post-in silico deformation geometry reduces peak contact pressure from about 27 MPa to about 14 MPa and yields more uniform contact stress distribution under near-full-extension running loads.

48. The device of claim 6, wherein the articular surface of the first component and the articular surface of the second component are prepared to receive heparin-conjugated hydrogel coating configured to promote lubricin adsorption for boundary lubrication upon exposure to synovial fluids.

49. The device of claim 41, wherein the cartilage domain polymer blend comprises a PLA / PBAT blend selected to tune Young’s modulus and yield stress for the articular region 50. The device of claim 49, wherein the blend comprises between about 55 / 45 and about 75 / 25 PLA / PBAT by weight.Docket 88800730-00052551. The device of claim 41, wherein the cartilage domain polymer and hydrogel are printed using a modified dual-printing unit comprising a Y-shaped nozzle and stepper motor-controlled feed to achieve homogeneous coating thickness and controlled infiltration depth.

52. The device of claim 1, wherein the first component and the second component are configured for press-fit implantation with an added slab in the first component to accommodate the multilayer build.

53. The device of claim 52, further comprising a custom guide to direct a first cut in a bone of a patient.

54. The device of claim 1, wherein the bone domain polymer-ceramic composite comprises PCL or PLA / PBAT with about 10 wt.% ZnO and about 20 wt.% HA / TCP, printed to achieve pore sizes of about 200-400 pm at about 6.5 mm by 6 mm by 1.2 mm dimensions for test coupons demonstrating cytocompatibility and ceramic loading fidelity.

55. The device of claim 39, wherein the cartilage layer coating is formed using hydrogel deposition at a nozzle height of between about 0.5 mm and about 1.0 mm above the scaffold surface to produce a uniform thin layer with controlled thickness by stepper motor speed.

56. The device of claim 30, wherein domain-specific target properties are specified to ensure compressive yield stress exceeding about 25 MPa in both bone and cartilage regions and bonecartilage tensile ultimate strength exceeding about 10 MPa.

57. The device of claim 31, wherein the scaffold features are configured to be manufactured with water-dissolvable support materials to preserve smooth articular surfaces after support removal.

58. The device of claim 18, wherein the ZnO component is included to provide piezoelectric activity to promote bone formation during physiologic loading without external electrodes.

59. The device of claim 31, wherein the osteochondral construct is dimensioned and contoured based on CAD models derived from micro-CT scans of conventional non-knee implants and modified by in silico plastic deformation to achieve improved congruence and reduced stress.

60. The device of claim 1, wherein the cartilage layer and the bone layer are configured and validated by ex situ compressive loading up to about 10 body weights without gross damage and with load-displacement curves absent discontinuities indicating failure.

61. The device of claim 1, wherein the cartilage layer polymer is subjected to ethylene oxide sterilization without significant change in tensile or compressive properties prior to implantation.Docket 88800730-00052562. The device of claim 1, wherein an internal structure and a print fidelity of the first component and the second component are verified by micro-CT of triangular infill architectures assessing pore geometry, fiber width, pore height, and pore width.

63. The device of claim 1, wherein the device is compatible with ancillary bioabsorbable fixation devices used clinically for osteochondral allograft surgeries to provide initial compression and shear strength comparable to metallic fixation.

64. The device of claim 1, wherein the cartilage layer is engineered to recapitulate interstitial fluid pressurization and boundary lubrication mechanisms of native articular cartilage under sliding loads.

65. A 3D printing nozzle, comprising:a first channel; anda second channel, wherein the nozzle is configured for printing one or more hydrogels, cells, and biomaterials onto a first scaffold layer via the first channel to support the growth of bone and onto a second scaffold layer via the second channel to support the growth of cartilage.

66. A method for bioprinting single-material and dual-formulation hydrogels to generate a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising:providing a bioprinter comprising a modified dual-channel nozzle configured to dispense hydrogel formulations under controlled pressure and feed rate;preparing a first hydrogel formulation comprising HAMA or fibrinogen and thrombin crosslinked with genipin to form a fibrin-based hydrogel (FibGen);optionally preparing a second hydrogel formulation comprising HAMA or FibGen conjugated with heparin to promote lubricin adsorption for boundary lubrication;loading the first hydrogel formulation into a first channel of the nozzle and the second hydrogel formulation into a second channel of the nozzle;controlling dispensing parameters including feed rate and nozzle height to deposit the hydrogel formulations onto a porous polymer scaffold corresponding to an articular surface region of the component;selectively printing either:the first hydrogel formulation alone to form a single-material hydrogel coating; orDocket 88800730-000525both the first and second hydrogel formulations in a layered or blended configuration to form a dual-formulation hydrogel coating;infiltrating the hydrogel formulations into the scaffold pores to a predetermined depth to enhance mechanical integration and lubrication properties; andcuring the printed hydrogel coating to achieve a target compressive modulus suitable for cartilage-like load-bearing performance.

67. A method for layered bioprinting of hydrogels onto a multi-domain scaffold for osteochondral applications, the method comprising:providing a three-dimensional printed scaffold comprising a bone-supporting domain and a cartilage-supporting domain;depositing a hydrogel onto a first scaffold layer to support growth of bone; and depositing a hydrogel onto a second scaffold layer to support growth of cartilage, wherein the nozzle is configured to print hydrogels, cells, and biomaterials onto the respective scaffold layers.

68. A bioprinter ink nozzle for dispensing multi-component mixtures to generate a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the nozzle comprising:a proximal manifold configured to mount on a dual-syringe holder and receive first and second precursor streams respectively from a first syringe and a second syringe;a bifurcated internal flow path including:first and second inlet channels fluidically coupled to the proximal manifold; a Y-shaped junction having a bifurcation angle that merges the first and second inlet channels into a common outlet path to combine the precursor streams immediately prior to dispensing; anda terminal section comprising one of:a screw-type static mixing segment configured to promote homogenous blending while limiting premature gelation; ora two-channel co-exit segment configured to deliver laminar co-flow of the first and second streams without internal mixing,a distal nozzle tip having an axisymmetric outlet geometry sized to deposit a continuous hydrogel filament with a target layer thickness and infiltration depth into an underlying porous scaffold;Docket 88800730-000525a drive interface configured for micropreci se control of dispensing by a stepping- motor actuation of the syringes to set a feed rate independent of pneumatic -pressure;an adjustable standoff feature that sets a nozzle height relative to the scaffold surface to control the printed hydrogel layer thickness and infiltration;a surface energy and wetting control treatment on at least the distal nozzle tip to improve filament continuity and reduce bulge formation during -deposition; anda modular tip architecture permitting interchangeable terminal segments to switch between the screw type static mixing segment and the two- channel -coexit- segment, wherein the nozzle is configured to print single material formulations from the first syringe alone, dual- formulation -coflow from both syringes, or blended formulations produced within the Y— shaped junction or the screw- type- static mixing segment, so as to achieve homogenous hydrogel coating thickness and controlled infiltration depth on an articular surface region of the scaffold.

69. A method of bioprinting a multi-component hydrogel using a bifurcated nozzle, comprising:providing a first hydrogel solution and a second hydrogel solution to a Y-shaped nozzle having a first channel and a second channel;mixing the first and second solutions at or immediately prior to an outlet of the nozzle; dispensing the mixed solutions onto a scaffold surface while controlling the extrusion feed rate via a stepper motor and the relative motion of the printhead; andselecting a nozzle height above the scaffold to form a homogeneous thin hydrogel layer with controlled infiltration.

70. A computer-implemented method for in silico modeling of plastic deformation to attain an articular surface shape for a scaffold of a biologic implant, the process comprising:obtaining scan data of a joint implant component comprising at least of a first component, a second tray and polyethylene insert by micro-computed tomography (microCT), the scan data capturing articular and bone-interfacing geometries;constructing, from the scan data, surface models and a stereolithography (STL) file by segmenting and smoothing the scan data to reduce imaging artifacts and yield rendered surfaces;importing the STL file into a computer aided design (CAD) environment and generating CAD models of the first component and the second component, including modifying patient- bone- interfacing regions to define flat mating surfaces;Docket 88800730-000525generating a finite element analysis (FEA) mesh from the CAD models of the first component and the second component, the mesh defining distinct articular and bone scaffold -domains;performing FEA with an elastic perfectly- -plastic material model applied to at least the articular layer, and applying a load representative of activities of daily living at near full extension, the load comprising a compressive force up to about ten body weights [and, optionally, an internal / external rotational moment];inducing, in silico, plastic yielding over at least a portion of the articular layer to increase first-second congruence and reduce peak contact pressure, and saving a deformed geometry when the deformation approaches a target cartilage -thickness;defining an articular surface shape based on the deformed geometry, includingfor the first component, extruding the deformed articular surface to form an articular domain, andfor the second component, sectioning the CAD model at a desired average cartilage thickness and inclination and establishing tied contact across split domains; iteratively altering at least one of the second or the first articular geometry using results of the FEA to further reduce peak contact pressure below a yield stress threshold for a selected polymer blend, wherein one or more alterations result in a finalized articular surface shape; and exporting the finalized articular surface shape as manufacturing data for scaffold fabrication and subsequent mechanical validation.

71. A computer-implemented method for generating a patient-specific articular surface shape for a non-knee implant by in silico plastic deformation, the method comprising:acquiring anatomical geometry of at least one component of a conventional non-knee implant via micro-computed tomography;segmenting scan datasets to generate surface models notwithstanding metal-induced artifacts; converting the segmented surface representations into a CAD model;generating a finite element analysis (FEA) mesh from the CAD model;applying boundary conditions and loads representative of activities of daily living to the FEA mesh and simulating elastic-perfectly plastic deformation of at least one articular component to increase articular surface congruence;saving the deformed geometry as an updated articular surface; andDocket 88800730-000525exporting the updated articular surface as a file suitable for manufacturing.

72. A method for scaffold failure analysis of a biologic equivalent of a component of a conventional non-knee implant, the method comprising:providing a scaffold comprising an articular surface region and a bone-interfacing region, the scaffold fabricated using 3D printing based on a computer-aided design (CAD) model derived from anatomic scan data of a conventional non-knee implant component;capturing a baseline three-dimensional geometry of the scaffold using micro-computed tomography (microCT) to generate volumetric imaging data of the scaffold structure;subjecting the scaffold to a mechanical load representative of physiologic conditions, the load comprising a compressive force up to about ten body weights applied at near full extension;rescanning the scaffold using microCT after the mechanical load to generate post-load volumetric imaging data;comparing the baseline and post-load volumetric imaging data to identify structural changes indicative of failure, the structural changes comprising at least one of crack initiation, pore collapse, strut fracture, or permanent deformation of the articular surface region; and reporting the failure analysis results for validation of scaffold design and material selection.

73. A method for controlled scaffold manufacturing and processing using flexible, degradable biopolymers for biological use to generate a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising:selecting at least one flexible, degradable biopolymer from the group consisting of polylactic acid (PLA), polycaprolactone (PCL), polybutylene adipate terephthalate (PBAT), and blends thereof, optionally combined with ceramic fdlers to promote osteoconduction;preparing a printable filament or feedstock by melt blending the selected biopolymer or biopolymer blend under controlled temperature and shear conditions to achieve homogeneity and target filament diameter;processing the blended biopolymer into a filament or pellet form suitable for fused deposition modeling (FDM) or direct bioprinting;defining an appropriate multiplier for each polymer or feedstock to control amount of material extruded in 3D printing / FDM process using flexible biopolymers;generating a computer-aided design (CAD) model of a scaffold corresponding to an articular surface region and a bone-interfacing region of the component, the scaffold comprising aDocket 88800730-000525porous architecture with a predetermined infdl pattern and porosity to balance strength and cellular infiltration;printing the scaffold using the prepared biopolymer feedstock under controlled extrusion parameters including nozzle temperature, layer height, infill percentage, and multiplier to achieve dimensional fidelity and mechanical integrity;optionally incorporating a gradient in material composition or porosity between the articular region and the bone region to mimic native osteochondral structure;post-processing the printed scaffold to remove support material and smooth the articular surface, and conditioning the scaffold under physiologic conditions to verify degradation rate and mechanical performance; andvalidating the scaffold for biological use by confirming biocompatibility, degradability, and mechanical properties suitable for load-bearing in a non-knee joint.

74. A method for creating a ceramic containing layered biologic, comprising:solvent casting of ceramic nanoparticles evenly into solubilized polymer using mixing or sonication between about 10% to about 40% w / v;evaporation of excess solvent and creation of polymer-ceramic composite flakes; shredding and granulation of the polymer-ceramic composite flakes;evaluating the shredded and granulated polymer-ceramic composite for melt flow index for flowability in melt blending;melt blending of the polymer-ceramic composite and extrusion of polymer-ceramic into filament for use in 3D printing and;3D printing, Fused deposition modeling, or direct printing of polymer-ceramic composite feedstock into controlled layers, scaffolds, and structures.

75. A layered biologic non-knee implant, comprising:an anatomically shaped scaffold body printed in a single, continuous three-dimensional (3D) build without post-print assembly seams, the scaffold body defined by a computer-aided design (CAD) model derived from scan data of a conventional non-knee implant component; a cartilage layer forming an articular surface region of the scaffold body, the cartilage layer comprising a flexible, degradable polymer or polymer blend and a hydrogel coating comprising a genipin-crosslinked fibrin hydrogel configured for boundary lubrication and interstitial fluid pressurization;Docket 88800730-000525a bone layer forming a bone-interfacing region of the scaffold body, the bone layer comprising a polymer-ceramic composite selected to promote osteoconduction and osteointegration;a printed boundary layer that is co-formed during the single, continuous 3D print to transition between the cartilage layer and the bone layer, the boundary layer comprising a gradation in at least one of material composition, porosity, or infdl orientation to maintain mechanical continuity and enhance cellular communication across layers;an internal porous architecture within at least one of the cartilage layer, boundary layer or bone layer, having a predetermined infill pattern and porosity selected to balance load-bearing strength and hydrogel infiltration depth; anda surface-smoothing feature achieved by printing with a water-soluble support material that is removed post-print to reduce micro-abrasions on the articular surface region, wherein the anatomically shaped scaffold body is configured to reduce peak contact pressure under physiologic loading and to provide layer-specific mechanical properties suitable for cartilage-like tribology at the articular surface and for osteoconductive integration at the bone-interfacing region.

76. A method for preparing a scaffold for a biologic equivalent of a component of a conventional non-knee implant, the method comprising:providing a three-dimensional (3D) printed scaffold comprising at least one of a first component or a second component, the scaffold including a bone-interfacing region and an articular surface region, the scaffold formed from a biocompatible and bioresorbable polymer or polymer-ceramic composite;subjecting at least a portion of the scaffold to plasma treatment under controlled pressure and exposure time to modify surface energy and enhance cell adhesion and hydrogel infiltration;rinsing or conditioning the plasma-treated scaffold in ethanol (ETOH) to remove residual contaminants and to sterilize the scaffold prior to biological use;optionally repeating the plasma treatment and ethanol conditioning steps to achieve a predetermined surface wettability and sterility level;sterilizing the scaffold using ethylene oxide;drying the scaffold under aseptic conditions and verifying that the scaffold maintains dimensional fidelity and porosity after plasma and ethanol processing; andDocket 88800730-000525releasing the scaffold for subsequent cell seeding or hydrogel infusion upon confirmation of sterility and surface activation.

77. A quality-control method for compositional analysis of 3D-printed polymer or polymerceramic filaments and detection of contamination, comprising:acquiring attenuated total reflectance Fourier transform infrared (ATR-FTIR) spectra of raw materials, intermediate pellets, and printed filaments;preprocessing the spectra by baseline correction and normalization;extracting spectral features corresponding to polymer and ceramic constituents, including PLA and PBAT band signatures and phosphate peaks associated with 0-TCP and HA;comparing extracted features to a reference spectral library derived from qualified inputs and prior acceptable prints to (i) quantify composition and verify no unanticipated changes due to solvent processing, melt blending, or printing and (ii) flag unexpected peaks or out-of-tolerance residuals as contamination; andoutputting a pass / fail decision and recommended corrective actions for the filament production lot.

78. A method for introducing cell suspensions to inoculate cylindrically shaped scaffold constructs for a biologic equivalent of a component of a conventional non-knee implant, the method comprising:providing a three-dimensional cylindrically shaped scaffold corresponding to at least one of a first component or a second component, the scaffold comprising a flexible, degradable polymer or polymer-ceramic composite with an internal porous architecture;preparing a cell suspension comprising autologous adipose-derived stem cells (ASCs) and / or allogeneic human induced pluripotent stem cell (hiPSC)-derived chondrogenic or osteogenic cells;mixing the cell suspension with a hydrogel carrier comprising HAMA or a genipin-crosslinked fibrin hydrogel (FibGen), optionally heparin-conjugated to promote lubricin adsorption and boundary lubrication;loading the cell -containing hydrogel into a bioprinter having a modified dual-channel nozzle and dispensing the mixture under controlled feed rate and nozzle height to deposit onto the scaffold surface;Docket 88800730-000525inoculating the cylindrically shaped porous construct by axial deposition and controlled infdtration of the cell-containing hydrogel to a predetermined depth within the pores;inoculating the cylindrically shaped scaffold by conformal deposition of the cell-containing hydrogel over an articular surface region and, optionally, into a bone-interfacing region, while maintaining a continuous coating thickness;controlling coating thickness and infiltration depth by adjusting the dispensing feed rate and nozzle standoff height; andverifying post-inoculation distribution and viability by imaging and mechanical adequacy of the coating, and releasing the construct for subsequent culture or implantation.

79. A scaffold for a biologic equivalent of a component of a conventional non-knee implant, comprising:an articular surface region configured to present a low-friction interface, the articular surface region including:a hydrogel coating comprising HAMA or a genipin-crosslinked fibrin hydrogel (FibGen) conjugated with heparin to promote boundary lubrication by lubricin adsorption; anda controlled infiltration of the hydrogel into a porous polymer scaffold underlying the articular surface region, wherein the scaffold further comprises:a surface-smoothing treatment configured to reduce micro-abrasions on the articular surface region resulting from fabrication supports; anda process control of hydrogel application comprising at least one of a defined nozzle height and a defined feed rate to achieve a target hydrogel layer thickness and infiltration depth, and wherein, when the scaffold is immersed in synovial-like fluid, the hydrogel coating provides interstitial fluid pressurization that bears a majority of an applied compressive load and lowers an effective friction coefficient at the articular surface region.

80. An articular-surface-treated non-knee scaffold comprising:a porous polymer or polymer-blend lattice; anda hydrogel coating on an articular region, wherein the coating comprises genipin-crosslinked fibrin formulated to promote lubricin adhesion for boundary lubrication and is deposited as a homogeneous thin layer tuned by feed rate and nozzle height.

81. A patient-specific implantation guide and fixation system for a layered biologic equivalent of a component of a conventional non-knee implant, comprising:Docket 88800730-000525a custom surgical guide body defined by a computer-aided design (CAD) model generated from scan-derived geometry of the conventional non-knee implant component and corresponding patient bone interface regions;one or more guide features that constrain bone resection to produce flat mating surfaces at the first and / or second interfaces consistent with the biologic implant’s scaffold design;thickness control features that reference an articular layer and a bone scaffold layer of the biologic implant and define cut planes or stop surfaces so that the prepared host bone accommodates a target cartilage-layer thickness and a complementary bone-layer seating depth;an output comprising digital tool-paths or positional data compatible with robotic or computer-assisted surgery for placement of the biologic implant in a single, continuous 3D-printed form without post-print assembly seams; anda fixation technique configured for non-permanent fixation, comprising press-fit seating of the biologic implant’s porous scaffold against the prepared flat mating surfaces to promote osteointegration, wherein the custom surgical guide body, the thickness control features, and the robotic or computer-assisted placement collectively align the articular surface region to reduce peak contact pressure under physiologic loading, and the fixation technique avoids cemented or permanent hardware while enabling osteoconductive integration of the bone scaffold layer.

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