Scaffold free cartilage implants

Abstract

The invention relates to methods of preparing an implant for bone repair, the method comprising the steps of expanding periosteum derived cells, mixing the expanded periosteum derived cells with a partially cross-linked sodium alginate solution, forming the alginate solution and the cells into a shaped construct, differentiating the cells in the shaped construct towards chondrogenic differentiation, and dissolving crosslinked alginate with a calcium chelating agent.

Description

[0001] SCAFFOLD FREE CARTILAGE IMPLANTS

[0002] FIELD OF THE INVENTION

[0003] The invention relates to the bioprinting of cells.

[0004] BACKGROUND OF THE INVENTION

[0005] Next generation Tissue Engineered Advance Therapeutic Medicinal Products (TE- ATMPs aim to regenerate defects and restoration of damaged tissues and organs TE- ATMPs implants, to a certain degree, must possess the ability to recapitulate the structure of the target host tissue with the correct cellular composition and functional properties. TE-ATMP solutions have become necessary for severe clinical conditions including skeletal critical size bone defects. To tackle these hurdles bone graft substitutes, such as hydroxyapatite, tricalcium phosphate, bioactive glasses and synthetic materialized implants are trying to mimic bone properties and enhance defect regeneration However, for critical size long bones defects, defects which are larger than 2.5cm, there is currently no regenerative solution, so patients need to undergo lengthy surgical interventions as well as revision surgeries which results in longer hospitalization.

[0006] Long bone fractures healing occurs through the formation of a transitional fibrocartilaginous template known as the fracture callus which progressively mineralizes resulting into a fully remodeled bone. This cartilage to bone transition, known as endochondral ossification (ECO) is a well-regulated deterministic biological process of that is also encountered in embryonic long bone development. During the past decade there is a rapid growth of developmentally inspired tissue engineering strategies that aim to closely recapitulated this cartilage to bone transition process. This field termed as "developmental engineering" aims to closely recapitulate those robust biological events encountered in embryonic development into the formation of tissue engineered products able to regenerate long bone defects. Cellular condensations able to mimic the mesenchymal condensation, a pivotal event in limb development, have been achieved in vitro through various setups such as cells spheroids and organoids- cell sheets- or using hydrogel carriers [Teng et al. (2022) Gels 8, 275] and progenitor cell types such as bone marrow MSCs- adipose derived MSCs and human periost derived progenitors (hPDCs) These formats have enabled the engineered of tissue engineered cartilaginous building blocks that are able to form bone upon implantation. However, for clinical use larger implants need to be produced. Bottom-up tissue engineering strategies enables the biofabrication of implants based on the modular assembly of these building blocks, for bone tissue regeneration upon implantation in small animal models, following developmental engineering paradigm [Ouyang et al. (2020) Adv. Fund. Mater. 30, 1909009]. However, organoid bioassemblies are formed through self-assembly processes and hence large-scale and anatomically complex cartilaginous implants with geometric precision cannot be produced. Moreover, technological limitations in scaling these implants towards dimensions matching those of critical size long bone defects are preventing further translation of these strategies.

[0007] A critical bottleneck towards clinical implementation is the ability to manufacture such implants at a clinically relevant scale and affordable cost. Bottom-up tissue engineering strategies enable the biofabrication of implants based on the modulaR. assembly of living building blocks [Liu et al. (2017) Bone Res. 5, 17014; Hernandez- Gonzalez (2020) Carbohydrate Pol. 229, 115514]. Bone-forming organoids have the capacity to self-assembly into larger bioassemblies, thus enabling a modular scalability [Gopinathan & Noh (2018) Biomaterials Research 22, S40824-018-0122-; Farshidfar et ai. (2023) Marine Drugs 21, 189] This bottom-up strategy faces technological limitations such as a limited scalability towards clinically relevant dimensions, costly and complex manual multi-step production processes, and a limitation regarding resolution of anatomical complexity and geometric precision resulting from a spontaneous aggregation process.

[0008] 3D Bioprinting proposes the combination of living cells or cell aggregates in rheological competent soft polymers (bioinks), sometimes enriched with bioactive molecules [Groll et ai. (2016) Biofab. 8, 013001]. Utilizing a layer-by-layer deposition of cellladen bioinks, tissues with increased shape complexity and dimensions can be produced However, bioprinting of bioinks into complex 3D structures demand the combination with different synthetic polymers which typically require multiple crosslinking steps in order for bioinks to solidify [Gopinathan, cited above]. This resulting in lengthy bioprinting processes where cells are exposed to ambient conditions while repeatedly exposed to cross-linking processes and can be harmful for the cells compromising their viability and capacity to differentiate post-printing, which is related to mechanotransduction Additionally, the presence of foreign material has been seen to affect the regeneration process and elicit immune responses from the host so solutions circumventing this challenge are further explored. Variety of hydrogels exploited for the development of bone and cartilage tissue engineering applications such as chitosan, collagen or gelatin, alginate, hyaluronic acid, heparin and poly(vinyl alcohol).

[0009] However, the presence of bioinks and exogenous bio-inks in implantable TE-ATMPs also includes risks such as immune response, lack of integration, limited efficacy as well as negative systemic effects of degradation Hence, the use of sacrificial bioinks is becoming an attractive strategy to retain the advantages of 3D-bioprinting without the associated risks. According to its characteristics including gelling capacity, low toxicity, biocompatibility, and cost efficiency, alginate hydrogels are extensively tested in bioprinting for bone tissue engineering (BTE) [Hernandez-Gonzalez, cited above; Farshidar et al. cited above]. Challenging limitation of alginate hydrogels for the biofabrication of complex 3D structures is the limited printability and shape fidelity even if concentration can be tuned Recently, novel treatment techniques have been established based on pre -or- partially crosslinking alginates prior to bioprinting improving the shape fidelity and the post bioprinting structural integrity of alginate bioinks [Hazur et al. (2020). Biofabrication 12, 045004; Falcone et al. (2022) Carbohydr Poiym 7, 100524].

[0010] SUMMARY OF THE INVENTION

[0011] In recent years endochondral ossification (ECO) became robust paradigm for bone tissue engineering solutions. This cartilage to bone tissue transition is encountered during bone development and fracture healing. ECO Tissue engineered solutions have been provided through high-density hydrogel-based strategies as well as organoid bioassemblies engineered cartilage templates exhibiting positive in vivo outcomes. However, these implants are mostly formed through self-assembly processes and hence large-scale and anatomically complex cartilaginous implants with geometric precision cannot be produced.

[0012] Moreover, the presence of foreign materials such as bioinks is well known to trigger immune response from the host. Therefore, a high cell-density bioprinting strategy, was developed using sacrificial alginate bioinks resulting in functional scaffold-free cartilaginous tissues able to rapidly form large and pure ossicles upon implantation. A chemically modified partially crosslinked alginate bioink was developed that could be printed in one extrusion step maintaining its structural form, using the advantage to 3D bioprint high cell densities. This bioink enabled the bioprinting of cartilaginous implants that could maintain structural precision and enable complex shapes to be produced. Upon chondrogenic differentiation of hPDCs within the bioink, condensed into clusters which gradually grew while abundant extracellular matrix was produced throughout the entire volume of the engineered tissues. Heedemonstrated that the alginate bioink used could be selectively dissolved at various timepoints resulting in structurally competent constructs. Finally, constructs were implanted ectopically to evaluate their bone formation capacity for 4 and 8 weeks.

[0013] The invention is further summarized in the following statements:

[0014] 1. A method of preparing an implant for bone repair, the method comprising the steps of: a) expanding periosteum derived cells for a period of between 7 and 28 days, b) mixing the expanded periosteum derived cells with a partially cross-linked sodium alginate solution, b) forming the alginate solution and the cells into a shaped construct, c) differentiating the cells in the shaped construct towards chondrogenic differentiation for a period of between 14 and 28 days, d) dissolving crosslinked alginate with a calcium chelating agent.

[0015] Herein the partially cross-linked sodium alginate solution is of such viscosity that one the one hand the solution is liquid enough for a bioprinting apparatus and on the other hand of a sufficient viscosity to maintain the shaped structure obtained upon bioprinting.

[0016] 2. The method according to statement 1, further comprising step e) of cultivating the cells after dissolving the cross-linked alginate for at least 5 days in the medium for chondrogenic differentiation, typically between 5 and 10 days, more typically for 7 days.

[0017] 3. The method according to statement 1 or 2, further comprising the step of after step b) and prior to step c) further crosslinking the obtained shaped construct with CaCL. This additional crosslinking of the shaped object provides a more rigid structure.

[0018] 4. The method according to statement 3, wherein the crosslinking is performed by placing the shaped construct in a 1-3 % CaCk (w / v) for between 1 and 5 minutes.

[0019] 5. The method according to statement 3, wherein the crosslinking is performed by placing the shaped construct in a 2% CaCk (w / v) for 2 minutes.

[0020] 6. The method according to any one of statements 1 to 6, wherein the partially crosslinked alginate in step b) has a tan(<5) between 0.1 and 0.15, at frequency of 1 Hz or between 0.12-0.13.

[0021] 7. The method according to any one of statements 1 to 6, wherein between 10 and 40 million cells are added per ml of alginate solution. The method according to any one of statements 1 to 7, wherein between 15 and 30 or between 20 and 25 million cells are added per ml of alginate solution. The method according to any one of statements 1 to 8, wherein the cross-linked alginate in step b) prepared by adding CaCI to a 4% (w / v) alginic acid sodium salt solution until a concentration of between 14 and 25 mM, typically between 17 and 21, more typically 19 mM CaCI is obtained. The method according to any one of statements 1 to 9, wherein forming the alginate and the cells into a shaped construct is obtained by bioprinting. The method according to any one of statements 1 to 11, wherein expanding periosteum derived cells in step a) is performed for a period of between 18 and 24 days, typically for 21 days. The method according to any one of statements 1 to 12, wherein differentiating the cells in step c) is performed for a period of between 18 and 24 days, typically for 21 days. The method according to any one of statements 1 to 12, wherein in the step a) the cells are expanded in a serum comprising basic cell culture medium. The method according to statement 13, wherein the medium comprises 10% foetal bovine serum, antibiotics and antimycotics, the medium being free from growth factors. The method according to any one of statements 1 to 14, wherein in step c) the cells are differentiated in a serum free cell culture medium comprising Bone Morphogenetic Protein 2 (BMP-2), Growth Differentiation Factor 5 (GDF5), Transforming Growth Factor beta 1 (TGF-beta 1), Bone Morphogenetic Protein 6 (BMP-6), and basic fibroblast growth factor 2 (basic FGF-2). The method according to statement 15, further comprising antibiotics, ascorbate- 2-phosphate, Proline, Insulin, Transferrin, Selenium, Dexamethasone, and Rho inhibitor, such as Y-27632. The method according to any one of statements 1 to 17, wherein the chelator is EDTA. The method according to any one of statements 1 to 19, wherein the shaped construct is incubated in a medium comprising between 10 and 100 mM EDTA. An implant prepared by the method of any one of statements 1 to 18, for the treatment of a bone defect. DETAILED DESCRIPTION

[0022] Figures

[0023] Figure 1. Synthesis and characterization of partially crosslinked alginate, a) printing resolution measurements. Resolution and accuracy analysis of three different partially crosslinked degree lOmM CaCk, 19mM CaCk, 26mM CaCk. b) Rheological analysis, c) Complex 3D printed constructs, d) Complex 3D bioprinted constructs.

[0024] Figure 2. Partially crosslinked alginate allows cell condensation and aggregation due to high single cell density a) brightfield and live / dead staining, b) interactive 3d surface plot of lived dead analysis.

[0025] Figure 3. Partially crosslinked alginate allows cell condensation and aggregation due to high single cell density) a) brightfield images over the differentiation timepoints, b) Quantification analysis of brightfield intensity, normalized with Day 0. c) Quantification analysis of aggregation area within the samples at Day 21. d) Live / dead staining over the differentiation timepoints, e) Quantification analysis of viability at the final differentiation timepoint Scale bars: B) 1mm, E) 200pm.

[0026] Figure 4. Evaluation of chondrogenic differentiation, a) Quantification of mRNA for chondrogenic and pre-hypertrophic / hypertrophic gene markers, b) metabolic activity. Figure 5. Evaluation of chondrogenic differentiation, a) Quantification of mRNA for chondrogenic and pre-hypertrophic / hypertrophic gene markers, b) immunostaining of COL2A1 cartilage matrix marker, c) Gradual increase in young modulus over the timepoints, related to abundant cartilaginous ECM secretion, d) representative stressstrain curve for the tested conditions e) Safranin 0 staining for cartilage-specific sulfated glycosaminoglycans (GAGs). f) Quantification of Alginate percentage at the final differentiation timepoint. Scale bars b) 1mm, e) 200um.

[0027] Figure 6. Gradual increase in young modulus over the timepoints, related to abundant cartilaginous ECM secretion, a) Safranin 0 staining black arrow represent partially crosslinked alginate compartments, red arrows represent cartilaginous extracellular matrix deposition, b) alcian blue staining, c) confocal z-projection dapi / phalloid in, organized collagen (2H). d) compression mechanical characterization.

[0028] Figure 7. Partially bone formation capacity upon ectopic implantation in small animal models for PC_ALG constructs, a) 3D rendering of nano-CT upon 4 and 8 weeks ectopic implantation, b) quantification of mineralized tissue / explant volume. c,d) Safranin 0, Trichrome Masson staining for 4 and 8 weeks.

[0029] Figure 8. Partially crosslinked alginate dissolution resulting to scaffold free standalone construct (SF 21D). a) FTIR analysis, b) live dead evaluation post and prior dissolution, c) brightfield imaging post and prior dissolution. d,e) Safranin O, Alcian blue staining for SF_21D and SF_28D.

[0030] Figure 9. Partially crosslinked alginate dissolution resulting to scaffold free standalone construct (SF-Day 21). a) Safranin O staining for ALG-Day 0, ALG-Day 21, SF- Day 21 and SF-Day 28. b) Fourier-transform infrared (FTIR) spectroscopy analysis, c) live dead and brightfield imaging post and prior dissolution.

[0031] Figure 10. Scaffold-free 3D constructs show full bone formation upon in vivo implantation (SF_21D, SF_28D). a) 3D rendering of nano-CT upon 4- and 8-weeks ectopic implantation, b) quantification of mineralized tissue / explant volume, c, d, e, f) Safranin 0, Trichrome Masson staining for 4 and 8 weeks.

[0032] Figure 11. Human size implant aa) Live dead staining. b,c) Safranin O and Alcian Blue staining.

[0033] The present invention discloses a single step high single-cell-density 3D bioprinting process educing a partially cross linked sacrificial bioink resulting in functional scaffold-free biologically functional and mechanically competent cartilage implants. Despite a high cell density high survival was observed and a cartilaginous phenotype was seen upon chondrogenic induction of human Periosteum Derived cells (hPDCs) within the partially crosslinked alginate bioink. Cells initially condensed into cell aggregates secreting abundant cartilaginous extracellular matrix throughout the entire volume of the engineered tissues. Partially crosslinked alginate bioink could be selectively dissolved at the end of the differentiation process, resulting in 3D scaffold- free stand-alone structurally competent constructs. Bone formation capacity and bone maturation were significantly improved for 3D scaffold-free stand-alone implants compared to partially crosslinked alginate implants. Pure ossicles exhibiting cortical and trabecular bone structure while a fully developed bone marrow cavity was observed upon ectopic in vivo implantation in small animal models. Surprisingly post dissolution the samples were mechanically and biologically robust stand-alone implants, although no references were found from the literature for stand-alone scaffold-free 3D bioprinted constructs. Moreover, this demonstrated biofabrication assay, paving the way for GMP-grade 3D bioprinted scaffold-free-stand-alone implants to address current clinical limitations.

[0034] Towards to clinical translational needs, Scaffold free TE therapies have been demonstrated. Relying to cell secrete ECM (extracellular matrix), cell sheets, have be revolutionized taking advantage the close cell-cell and cell-ecm interactions for the biofabrication of autonomously engineer microtissues. Current limitation possess that a 3D scaffold free tissue is that require rich-ECM content resulting to high number of cells. Limitation of scaffold free TE is that in order to fabricate a larger 3D scaffold multiple cell sheets need to be stacked, resulting to cell necrosis at the middle layer34. To address this dimensionality limitation scaffold free systems have been combined with graft substitutes like hydrogels and particles.

[0035] The present invention discloses a bioprinting-based biofabrication process able to produce scaffold-free stand-alone cartilaginous tissue with dimensions relevant to clinical practice. This process made use of a partially crosslinked linked bioink seeded with high density of human periost derived cells enabling their condensation and differentiation followed by the secretion of abundant cartilaginous extra cellular matrix. Periostel cells contain osteo- & chondro- progenitor cells and are key players in the formation of the fracture callus and the main contributors to the process of endochondral ossification. The removal of the periosteal layers has been shown to hamper long bone defect healing. Therefore, they constitute the ideal progenitor cell type when aiming to heal large long-bone defects.

[0036] As bioink alginate is selected due to its inert nature and due to the fact that it does not provide binding sites for the cultured cells forcing them to condense and selfaggregate. By partially crosslinking alginate, a novel bioprintable material was developed able to be extruded through a bioprinter nozzle retaining the shape defined by the bioprinting process without any further cross linking steps. The presence of pre cross linked alginate ensured an initial structural support to the cells that over time lost its importance as the cells secreted mature collagen as evidenced by second harmonics imaging. Therefore, as a follow up alginate was sacrificed / dissolved and the chondrogenic differentiation process was continued in the absence of any foreign material. The resulting tissue was seen to possess a striking bone forming potential resulting in the formation of pure ossicles after 4 and 8 weeks of implantation. During explant analysis, the bioprinted implants exhibited cortical and trabecular bone structures as well as abundant bone marrow compartment.

[0037] In order to scale up the implant to dimensions relevant to critical size defects scale up in both the production of single cells as well as scale up of the bioprinting process is required. Therefore, a clinical grade bioreactor was employed to expand hPDCs seeded on microcarriers in suspension. This process delivered the same amount of cells as expected for treating critical size tibial defects. Once the cells were harvested from the bioreactor a cylindrical ring tissue engineered construct was bioprinted and the construct was placed within a perfusion bioreactor for the subsequent chondrogenic differentiation process. This allowed to recapitulate observations made in smaller scale constructs and strikingly allowed to dissolve the alginate compartment leading to a 2.5 cm implant able to possess structural integrity composed of only the cells and the secreted cartilaginous matrix. Characterization of this implant revealed homogeneity of the secreted extracellular matrix and the presence of high fraction of living cells.

[0038] The present invention combines the use of sacrificial partially crosslinked alginate bioink with high-density loading of extracellular matrix producing cells, to create large autologous scaffold-free bone forming implants. Specifically, a biologically inert alginate bioink was chemically modified for improved printability allowing high cellular density. Then, cells were stimulated towards chondrogenic differentiation and cartilage ECM production. Towards chondrogenic differentiation, the retaining partially crosslinked alginate compartments within the samples were dissolved resulting in a scaffold free stand-alone tissue. The resulting implants were implanted ectopically in nude mice to assess their bone forming potency. Finaly, this one-step biofabrication strategy was used to produce a human-size stand-alone scaffold free autologous implant according to the dimensional specifications of a non-union bone defect.

[0039] Sodium alginate was employed as a bioink. However alginate does not possess the rheological properties to meet the extrudability standards for bioprinting a 3D implant. Sodium alginate was chemically modified with CaCh, resulting in a partially crosslinked alginate bioink with enhanced rheological properties, able to maintain its structural support with high fidelity and post printing accuracy . The rheological analysis outlined a significant increase in the storage (G') and loss (G") modulus of the partially crosslinked alginate, especially when comes to 19 mM CaCh of final CaCh concentration, which optimized conditions for bioprinting and structural integrity of the bioprinted constructs . The observed reduction in line width and the improved fidelity are consistent with previous studies that identified the role of accurately tuned viscoelastic properties in extrusion-based bioprinting. Increasement of viscosity is also supported by previous studies that described the beneficial impact of calcium ions on improving the printability and structural integrity of alginate-based bioinks. Due to the inert nature of alginate, and due to the fact that it does not provide cell binding sites, the cells spontaneously aggregated within the printed hydrogel. This cellular condensation is a deterministic first step in fracture callus formation and the initiation of chondrogenic differentiation. The cell and tissue fate was assessed during long term differentiation, as an effect of cell density, and thus post-printing aggregation capacity. The 3 tested conditions were compared upon chondrogenic differentiation for 21 days. Volumetric formation of cartilaginous microtissues exists exclusively at the highest tested concentration of hPD single cells (23 million cells / ml) within the sample followed by the secretion of abundant cartilaginous extracellular matrix, indicating higher chondrogenic differentiation potential, as evidenced by staining and gene expression analysis However, upon in vivo implantation of the selected condition (23 million hPDCs / ml), by pCT analysis was demonstrated partial bone regeneration and poor in vivo performance particularly at the periphery of the constructs as well as within the central region Figure 7.

[0040] Recapitulating the in vitro analysis, the present invention demonstrates that partially crosslinked alginate was actively replaced by autologous cell-produced ECM throughout the differentiation process, as the abundant ECM was considered superior to the partially crosslinked alginate compartments retaining within the constructs. Histology and normalized alginate degradation analysis demonstrates that only 25% of the initial partially crosslinked alginate was retained within the sample at the final differentiation timepoint Still, this 25% of bioink retained in the construct hampered the in vivo remodeling with varied ectopic bone ossicle quality.

[0041] The present invention discloses a scaffold free self-supporting construct fabrication of implants. The dissolution of the 25% of the retaining partially crosslinked alginate, allows the production of constructs consisting exclusively of their own cells and their secreted ECM (Figure 9). Similar scaffold-free approaches have been explored in tissue engineering, but most rely on either ECM scaffolds or devitalization tissues. Dissolving the remaining partially crosslinked alginate compartments with the use of Ethylenediaminetetraacetic acid (EDTA chelating agent), results in a scaffold-free 3D bioprinted implant. In contrast with previous studies, chelation agents were used for softening and partially degrading alginate compartments, since alginate does not represent an ideal matrix for culturing some other types of cells (iPSC) due to its bio inertness and in vivo limited biodegradation

[0042] However, post alginate dissolution an additional week of further differentiation was necessary to achieve the ultimate structural integrity such that the final scaffold-free construct retains its structural geometry, shape, and mechanical integrity. The resulting 3d bioprinted scaffold-free construct was seen to possess a striking boneforming potential resulting in the formation of pure big ossicles even upon 4 weeks of implantation in small animal models. During the explant's analysis, the scaffold-free 3d bioprinted implants exhibited cortical and trabecular bone structures as well as abundant bone marrow compartments (Figure 10), demonstrating huge differences between the 3D bioprinted ALG-day 21 and scaffold-free explants.

[0043] For a better understanding of the mechanism by which partially crosslinked alginate removal influences the phenotype of the final implant and the significant enhanced in vivo performance transcriptomic analysis was carried out. A significant upregulation in the expression of critical early chondrogenic, late (pre)hypertrophic and osteogenic differentiation markers were demonstrated through transcriptomic analysis even 1 hour after alginate dissolution Meanwhile, a stable phenotype was observed in cells within the alginate, demonstrated by the overlapping PCA plots at day 21 and day 28 Post alginate dissolution, a shift backward to day 14 and continuous extreme upregulation after day 28 of the differentiation time point was observed. Furthermore, downregulation of immune modulatory genes such as IL2R. and CD2, along with significantly higher expression of genes related to extracellular matrix (ECM) production, was demonstrated following alginate dissolution. Overall, the samples that undergo alginate dissolution at the final timepoint (day 28) had a significantly higher expression of genes related to extracellular matrix (ECM) production, as well as differentiation toward the chondrogenic, connective tissue and osteogenic lineage, which, together are the main cell types in the fracture callus

[0044] EDTA is a chelating agent that binds to divalent cations such as calcium and magnesium dissolving that way the remaining alginate but also interfering with cell adhesion transmembrane receptors such as integrins as well as bindings of the matrisome associated proteins. Structural ECM proteins like fibronectin (cell adhesion, embryonic cell migration, wound repair) also bind to integrins. Ca2+removal might interfere with integrin signaling, and upregulated differentiation pathways and ECM production as fibronectin (FN1) were observed. Osteopontin (SPP1) and Bone Sialoprotein (BSP) matricellular proteins are also interacting with ECM through Ca-i- ions. Thus, post dissolution gene expression is upregulated because of the protein removal. This is a typical cellular feedback system. Presumably, proteins that adsorbed to the surface form a provisional matrix, influence the subsequent immune cell adhesion and activation witch found to be downregulated post dissolution. In particular, changing the surface chemistry will influence surface's protein adsorption, dictating the type of immune cell that will interact with it. This is supported by the proteomics analysis, demonstrating that higher number of proteins found post the dissolution presses. Further, the scalability shown in the present invention via GMP-compliant bioreactor systems addresses one of the key translational challenges noted in previous studies regarding the clinical-scale production of hPDC's, while is focused on scaffold-free implants towards to clinically relevant long bone defect sizes. To scale up the implant to dimensions relevant to critical size defects scale-up in both the production of single cells as well as scale-up of the bioprinting process is required. A clinical-grade bioreactor was used to expand hPDCs seeded on microcarriers in suspension. This process delivered the same number of cells as expected for treating critical-size tibial defects. Once the cells were harvested from the bioreactor a cylindrical ring tissue- engineered construct was bioprinted and the construct was placed within a perfusion bioreactor for the subsequent chondrogenic differentiation process (Figurell). It was possible to dissolve the alginate compartment leading to a 2.5 cm implant able to possess structural integrity composed of only the cells and the secreted cartilaginous matrix . Characterization of this implant revealed homogeneity of the secreted extracellular matrix and the presence of a high fraction of living cells (Figure 11 b,c). The present invention discloses a novel one-step biofabrication process for large, mechanically robust and biologically performant implants for bone regeneration that enabled the functional differentiation of hPDC single cells into a geometrically complex and scaffold-free bone forming implant. To conclude, this bioprinting process is efficient and scalable as demonstrated by the large human size scaffold-free cartilage implant and will support the translation of skeletal TE-ATMPs in clinical applications.

[0045] Example 1. Partially crosslinked alginate shows enhanced rheological properties for biofabrication of 3D complex structures.

[0046] Sodium alginate lacks the rheological properties to meet the extrudability standards in order to bioprint a 3d structure. To overcome this limitation, alginate was modified with calcium chloride (CaCL) (dihydrate CaCl2+2H20) synthesizing a partially crosslinked sodium alginate ink with improved viscosity . Line width and spreading ratio measurements (Figure la) showed that the ink with about 19mM CaCI concentration, is suitable for complex 3D structures that require high resolution.

[0047] To optimize the final crosslinked degree, three varied concentrations of CaCl2+2H2O (10 mM, 19 mM, and 26 mM) were titrated into the alginate, and was printed. Rheological measurements were conducted for bare alginate ink as well as for lOmM, 19mM, 26mM partially crosslinked inks (Figure lb).

[0048] Rheological measurements were conducted for the bare alginate ink as well as for the chemically modified 10 mM, 19 mM, and 26 mM partially crosslinked inks A significant change to the storage (G') and loss (G") modulus was observed, especially for the 19 mM and 26 mM inks, when compared to the lowest chemically modified ink (10 mM) In parallel, higher relative viscosity was observed for the 19 mM and 26 mM inks.

[0049] In parallel was observed higher relative viscosity for 19mM and 26mM inks. Partially crosslinked alginate (16mM) was loaded to the bioprinter and complex 3D structures were printed in larger scale (Figure lc). Post printing the contracts show structural fidelity while were printed without support or postprocessing, in one extrusion step maintaining its structural form. Next step was the mixing of partially crosslinked alginate (19mM) with single cell suspension of hPDCs in concentration of 23 million cell / ml. Large 3d complex structures were successfully 3D bioprinted, while post printing cell viability analysis were caried out (Figure Id). Cells exhibited high viability while the shape and structural stability were not significantly influenced by the high cell density.

[0050] Example 2. Partially crosslinked alginate allows cell condensation and aggregation due to high single cell density

[0051] Following the assessment of bioprinting process, structures underwent chondrogenic differentiation for 21 days. Over the timepoints of differentiation, cells demonstrated high viability while morphological differences were observed. Due to the high single cell density cells aggregate within the bioink forming microaggregates while progressively fuse into a larger condensated unified structures after 21 days (Figure 2a).

[0052] Example 3. Partially crosslinked alginate allows cell condensation and aggregation due to high single cell density (confirmation).

[0053] Human periosteum derived cells (hPDCs) were isolated from periosteal biopsies from various donors and were utilized as pooled cell samples. hPDCs were mixed with partially crosslinked alginate in various concentrations of 3 million, 13 million, and 23 million per milliliter and were 3D bioprinted in cylindrical constructs with dimensions 6mm in diameter and 2mm in height. Following the assessment of the bioprinting process, structures underwent chondrogenic differentiation for 21 days.

[0054] Morphology and viability of 3D bioprinted constructs were assessed at selected differentiation time points (day 0, 7, 14, 21). Morphological differences were observed macroscopically, while on the brightfield images, the intensity was higher over time only for 13 million cells / ml and 23 million cells / ml (Figure 3a). Following statistical analysis, brightfield normalized intensity (normalized to day 0) of the condition with 23 million cells / ml wass found to be statistically significantly higher (82%), while it possessed a higher aggregation area (54.63 ± 34 mm2) at the final differentiation timepoint (Day 21), compared to 13 million cells / ml, which was found to be 60% and 5.07 ± 2.5 mm2, respectively (Figure 3b, c). These results were probably indicating higher extracellular matrix secretion in the condition with 23 million cells / ml. The lowest tested concentration of 3 million cells / ml did not show differences regarding morphology, while the normalized intensity was quantified at 9.6%, and the aggregation area was 0.46 ± 0.15 mm2. On the other hand, spheroid formation due to the high single-cell density was observed only for the highest experimental conditions (13 million cells / ml and 23 million cells / ml), even from the 7th day of differentiation. Surprisingly, cells were aggregated within the bioink, forming microaggregates that progressively fused into larger, condensed, unified structures after 21 days of differentiation, only for the condition with 23 million cells / ml, in contrast with 3 million cells / ml, which showed a stable morphological structure from the 7th day until the end of the differentiation process.

[0055] Live / dead staining analysis demonstrated that all conditions achieved high viability over the differentiation time points (Figure 3d). Followed by quantification analysis, cell viability was maintained at 89-93% with no statistically significant differences after 21 days of differentiation for all the experimental conditions (Figure 3e).

[0056] Example 4. Evaluation of chondrogenic differentiation potential of the 3D Bioprinted constructs upon 21 days of culture

[0057] Fluorescent antibody staining was carried out for Indian hedgehog (IHH), matrix metalloproteinase-13 (MMP13), transcription factor osterix (OSX), proliferation marker (KI67), collagen type I (COL1A1), cartilage matrix marker (COL2A1) antibodies. Positive fluorescence antibody staining was observed for all the antibodies indicating the chondrogenic differentiation of the 3d bioprinted constructs. Gene expression of relevant markers was analyzed to further define the differentiation stages of the constructs (Figure 4a). Early chondrogenic transcription factor SOX9 was upregulated while the cartilage matrix marker collagen type II (COL2A1) was highly upregulated. The transcription factor osterix (OSX) which is directly regulated by (RUNX2) and expressed in pre-hypertrophic chondrocytes and osteoblasts was found to be upregulated over the timepoints, hypertrophic markers (COL10) and (IHH) were upregulated on day 21. Gene markers MMP13 VEGF and BSP indicating a pre-hypertrophic / hypertrophic phenotype were found to be upregulated. In addition, high metabolic activity was found on day 14 while glucose consumption increased, which is corresponding to the higher lactate values (Figure 4b).

[0058] Example 5 Evaluation of chondrogenic differentiation potential of the 3D Bioprinted constructs upon 21 days differentiation

[0059] Gene expression of relevant markers was analyzed to further define the differentiation stages of the constructs (Figure 5a). Cartilage matrix marker collagen type II, transcription factor osterix (OSX), which expressed in pre-hypertrophic chondrocytes and osteoblasts, were found to be upregulated over the time points, especially for the high single-cell concentration (23 million cells / ml) upon the final experimental time point. Hypertrophic marker (IHH) was found to be highly upregulated only for the condition with 23 million cells / ml on day 21.

[0060] Immunostaining of cartilage matrix marker collagen type II (COL2A1) was carried out for all the experimental conditions from Day 7 to Day 21 (Figure 5b). Positive staining was observed only for the 13 million cells / ml and 23 million cells / ml conditions after Day 14. When compared to the final time point, abundant cartilage collagen type II within the area of the construct was found only in the 23 million cell / ml condition. Alginate, being a natural polysaccharide, is intensively stained by Safranin O, which is used to stain cartilage-specific sulfated glycosaminoglycans (GAGs). Towards differentiation, higher secretion of extracellular matrix (ECM) can be observed, while alginate compartments within the samples were found to be less, especially for the 23 million cell / ml condition. In Figure 5f alginate degradation was quantified, showing that at the final time point, only 25% of the sample area corresponded to the presence of alginate for the condition of 23 million cells / ml, while 93% and 40% were found for the 3 million cells / ml and 13 million cells / ml conditions, respectively.

[0061] Mechanical characterization experiments were implemented to measure Young's modulus over the time points, to further define the extracellular matrix secretion toward chondrogenic differentiation (Figure 5c). Significant differences in Young's modulus were found even from the first time point due to the different cell concentrations within the constructs. Over the time points, only the 3 million cells / ml condition lost its mechanical properties, from 78.14 ± 9.8 kPa to 17.25 ± 5.1 kPa. At the final time point, 13 million cells / ml was found to have a slightly higher Young's modulus (76.68 ± 8.6 kPa) compared to the 23 million cells / ml condition (74.94 ± 14.4 kPa), while it possessed higher stress, as is demonstrated in Figure 5d. Example 6 Gradual increase in young modulus over the timepoints, related to cartilaginous ECM secretion

[0062] Alginate is a natural polysaccharide and is stained from Safranin O - Alcian blue staining, which are used for the stain of cartilage-specific sulfated glycosaminoglycans (GAGs). At (Figure 6a) black arrows indicate alginate compartments while red arrows the secreted ECM. Towards differentiation higher secretion of extracellular matrix (ECM) can be observed, linked to the constructs chondrogenic differentiation (Figure 6a, b). Filamentous-actin (F-actin) staining that revealed stress fibers in the actin cytoskeleton and second harmonics confocal analysis for organized collagen analysis were carried out (Figure 6c). Results show the present in abundance filamentous actin and organized collagen over the entire surface of the samples, upon 21 days of differentiation. Mechanical characterization experiments were implemented to measure young modulus over the timepoints (Figure 6d). Bare alginate is getting influenced from the differentiation media and loses approximately 100% of the young modulus after the 21 days, meaning that the alginate loses its mechanical properties over the differentiation presses. Hence, PC_ALG samples demonstrate over than 80% of young modulus increase, linked to abundant ECM secretion through the differentiation process as previously described.

[0063] Example 7. Partially bone formation capacity upon ectopic implantation in small animal models for PC_ALG constructs (Day 21 (23 million cell / ml) constructs.

[0064] Based on the in vitro results, the selected concentration to proceed with the in vivo studies was the 23 million cells / ml (ALG-Day 21), as a higher differentiation capacity was demonstrated compared to the other conditions.

[0065] PC_ALG constructs were implanted ectopically in immunodeficient nude mice to evaluate their capacity for bone formation upon implantation for 4 and 8 weeks.

[0066] Mineralization was detected and analyzed with Nano-Computed Tomography (Nano- CT) (Figure 9 a). Cortical and trabecular bone structures were formed, but not all of the construct area was mineralized even after 8 weeks upon implantation, as a nonmineralized cavity was located at the construct's center.

[0067] Quantification analysis was carried out for the Mineralized tissue / Bone volume (BV / TV %), and no statistical differences were found between the two in vivo timepoints (Figure 9 b). Bone compartments and bone maturation were assessed using Safranin O and Mason's trichrome staining (Figure 9 c,d). Bone marrow compartments were present at both timepoints, 4 and 8 weeks. Upon 8 weeks of implantation, an improvement in bone marrow compartments was observed.

[0068] Example 8. Partially crosslinked alginate dissolution resulting to scaffold free stand-alone construct (SF 21D)

[0069] This biofabrication strategy for the dissolution of the partially crosslinked alginate, results in a 3D tissue engineered construct composed of only "cells and their secreted extracellular matrix" and the absence of any foreign material. FTIR analysis was carried out to justify the absence of alginate at the SF free construct (Figure 8a). ECM derived from hPDCs microtissues was compared with PC_ALG and SF_21D lyophilized constructs, upon 21days of differentiation. Only the spectrum of PC_ALG shows the characteristic alginate at 1410 (cm'1) while SF_21D and ECM spectrums overlapping at collagen characteristic peaks (1654 / 1541 cm4). Live dead assay was used to evaluate the cell's viability prior and post dissolution (Figure 8b). The results did not show differences regarding the viability. Finally, it is important to mention that the 3D geometry of the initial 3D bioprinted construct (containing the bioink / hydrogel) is retained by the final scaffold-free Tissue Engineered construct alone, because of the high amount of abundant extracellular matrix that was secreted towards the differentiation process (Figure 8c). Post dissolution one additional week of chondrogenic differentiation were implemented, to enhance the mechanical integrity of the SF_21D construct. Safranin O and Alcian blue glycosaminoglycans (GAGs) histology stains for the samples SF_21D and SF_28D found to be positive for cartilage-specific extracellular matrix (Figure 8 d,e).

[0070] Example 9 Partially crosslinked alginate dissolution resulting to scaffold free stand-alone construct (SF-Day 21) [confirmation]

[0071] The dissolution of the partially crosslinked alginate results in a 3D tissue-engineered construct that is composed mainly of "cells and their secreted extracellular matrix," without any foreign material. At the final differentiation timepoint (Day 21), the differentiated samples (ALG-Day 21) are immersed in Ethylenediaminetetraacetic acid (EDTA) solution to dissolve the remaining alginate, resulting in scaffold-free, standalone tissue (SF-Day 21). It is important to note that the 3D geometry of the initial bioprinted construct (which contains alginate) is retained in the final scaffold-free tissue-engineered construct due to the substantial secretion of extracellular matrix throughout the differentiation process. Following dissolution, an additional week of chondrogenic differentiation is implemented to enhance the mechanical integrity of the SF-Day 21 construct . As described above, Safranin O intensively stains alginate compartments. Safranin O staining of glycosaminoglycans (GAGs) in the ALG-Day 0, ALG-Day 21, SF-Day 21, and SF-Day 28 samples reveals positive results for cartilage-specific extracellular matrix, while the absence of alginate is clearly observed in the SF-Day 21 sample post-dissolution (Figure 9a). Fourier-transform infrared (FTIR.) spectroscopy analysis was conducted to confirm the absence of alginate in the scaffold-free constructs. ECM derived from hPDCs microtissues is compared with lyophilized ALG-Day 21 and SF- Day 21 constructs. The characteristic alginate peak at 1410 cm- 1is observed only in the spectrum of ALG-Day 21, while SF-Day 21 and ECM spectra overlap at collagen characteristic peaks (1654 / 1541 cm- 1) (Figure 9b).

[0072] Finally, live / dead analysis shows that high cell viability is maintained following dissolution (Figure 9c).

[0073] Example 10. Scaffold-free 3D constructs show full bone formation upon in vivo implantation

[0074] Nano-CT analysis demonstrates fully developed big bone ossicles, consisting of cortical and trabecular bone structures even after 4 weeks upon implantation in immunodeficient nude for both experimental conditions SF_21D and SF_28D (Figure 10a). Following quantification analysis for Mineralized Tissue / Explant Volume (BV / TV %), statistically significant differences were found between the 8 weeks of implantation (Figure 10 a).

[0075] Safranin O and Masons trichrome staining were used to evaluate explants bone compartments and maturation. Mineralized cartilage compartments were found after 4 weeks of implantation for both experimental conditions (Figure 10 c,d). In contrast, upon 8 weeks of implantation there was no presence of mineralized cartilage especially for the SF_28D resulting to a fully bone formed tissue (Figure 10, f). Bone marrow cavities were fully developed for both implantation timepoints (Figure 10c,d,e,f).

[0076] Example 11. Human size implant

[0077] To address clinical translation needs, the previously described biofabrication strategy was demonstrated for the scaling-up of ATMPs. Donor-derived periosteum single cells were expanded from 1 million to 180 million over 30 days using a GMP-grade expansion bioreactor (SCINUS) The cell suspension was mixed with the partially crosslinked bioink and subsequently fabricated into two three-dimensional bioprinted constructs, with dimensions 2.5 cm in height and 2 cm in radius. Post-printing, the constructs were incubated for 21 days within a custom-designed perfusion bioreactor containing 250 mL of chondrogenic medium to provide necessary growth factors for cell differentiation .

[0078] On the 21st day, alginate dissolution was conducted, and chondrogenic differentiation was extended for an additional seven days. The total biofabrication time per bioprinted construct, including the crosslinking process, was recorded at 15 minutes.

[0079] Structural competency was demonstrated by the scaffold-free construct in maintaining its size and the original 3D bioprinted shape. Cell viability was assessed using a live-dead assay (Figure 11a).

[0080] Safranin O and Alcian blue staining were conducted to identify cartilaginous extracellular matrix in both experimental conditions (Figure 11 b). From the Safranin O staining, the absence of alginate was clearly observed in the SF-28 Day construct. Finally, the presence of cartilage-specific sulfated GAGs was demonstrated in both constructs after 21 days in culture.

[0081] Example 12 Material and Methods

[0082] Partially crosslinked alginate synthesis

[0083] Alginic acid sodium salt powder was dissolved in MiliQ ultrapure water (resistivity > 18.2 MQ*cm) in total concentration of 4% (w / v) and left at 4°C overnight to fully hydrate, under shaking conditions. Calcium chloride dihydrate solution (CaCl2+2H2O) was prepared in concentration of 90mM. To synthesize partially crosslinked sodium alginate, 1.5ml of 4% Alginic acid sodium salt was titrated with 0.4ml of 90mM CaCl2+2H2O with titration rate 0.02ml per minute, under 500rpm stringing speed and 40°C temperature. The final concentration of CaCL within the partially crosslink alginate ink was calculated 19mM.

[0084] Rheological analysis

[0085] A Discovery Hybrid Rheometer 2.0 (TA Instruments, DE, United States) was utilized with a 20mm parallel plate configuration. The geometry gap was consistently maintained at 200 mm for all bioink measurements. Initially, an oscillatory strain sweep ranging from 0.01% to 100% at an oscillation frequency of 1 Hz was carried out to identify the linear viscoelastic region (LVR). Subsequently, an oscillatory frequency sweep from 0.01 to 100 Hz with a strain fixed within the LVR was executed to monitor the variations in the storage modulus (G') and loss modulus (G"). Finally, the viscosity was assessed by incrementally increasing the shear rate from 0.01 to 100.

[0086] Human periosteum derived cells (hPDCs) expansion

[0087] Human periosteum derived cells (hPDCs) were isolated from periosteal biopsies from various donors and utilized as pooled cell samples. These hPDC pools were cultured until passage 7 at 37°C with 5% CO2 and 95% humidity in Dulbecco's modified Eagle medium (DMEM, Life Technologies, UK) supplemented with 10% fetal bovine serum (HyClone FBS, Thermo Scientific, USA) and 1% antibiotic-antimycotic (100 units / mL penicillin, 100 mg / mL streptomycin, and 0.25 mg / mL amphotericin B). The cells were subsequently used for various experiments (in vitro, in vivo, qPCR). All patients provided informed consent, and all procedures were approved by the ethical committee for Human Medical Research (KU Leuven) under approval number ML7861.

[0088] 3D Bioprinting process

[0089] Prior-bioprinting process, cells were trypsinized, counted and mixed with partially crosslinked bioink in a 5 mL Luer lock syringe at a final concentration of 23 million cells / mL. The syringe was then mounted onto the bioprinter cartridge (RegenHU). Using RegenHU research software (BioCAD), cylindrical constructs (6 mm diameter / 2 mm height) were printed in a 24-well plate. Post-bioprinting, the constructs were crosslinked with 2% (w / v) CaCL for 2 minutes, washed three times with Tris-HCI buffered saline (TBS) and transferred in 12-well non-adherent plate, while subsequently incubated in chondrogenic differentiation medium.

[0090] Chondrogenic differentiation

[0091] 3D bioprinted constructs undergo chondrogenic differentiation in a serum-free, chemically defined chondrogenic medium (CM) containing LG-DMEM (Gibco), supplemented with 1% antibiotic-antimycotic (100 units / mL penicillin, 100 mg / mL streptomycin, and 0.25 mg / mL amphotericin B), 1 mM ascorbate-2-phosphate, 100 nM dexamethasone, 40 pg / mL proline, 20 pM Rhokinase inhibitor Y27632 (Axon Medchem), ITS+ Premix Universal Culture Supplement (Corning) (including 6.25 pg / mL insulin, 6.25 pg / mL transferrin, 6.25 pg / mL selenious acid, 1.25 pg / mL bovine serum albumin (BSA), and 5.35 pg / mL linoleic acid), 100 ng / mL BMP-2 (INDUCTOS), 100 ng / mL growth / differentiation factor 5 (GDF5) (PeproTech), 10 ng / mL TGF-01 (PeproTech), 1 ng / mL BMP-6 (PeproTech), and 0.2 ng / mL basic FGF-2 (R&D Systems) Half of the medium was refreshed every 2 days. Viability assay

[0092] Light microscopy was employed at selected time points to evaluate cell growth. Additionally, cell viability was qualitatively assessed using the LIVE / DEAD Viability / Cytotoxicity Kit (Invitrogen, USA), following the manufacturer's protocol. Briefly, the 3D bioprinted constructs were rinsed with TBS, then incubated with 2 pM Calcein AM and 4 pM ethidium homodimer-1 for 30 minutes at 37°C with 5% CO2 and 95% humidity. The stained samples were visualized using a fluorescent Olympus LS microscope. qPCR

[0093] Constructs (n=3) collected at selected timepoints (Day 0,7,14,21), snap-freezed into a liquid nitrogen container and stored at -80 °C. Followed by cell lysis in 350 pL R.L.T lysis buffer (Qiagen, Germany) and 3.5 pL p-mercaptoethanol (Sigma Aldrich, Germany), samples were placed in 2.8 mm ceramic beads for hard tissue homogenizing. RIMA was extracted from the 3d bioprinted contracts using Nucleospin RNA II kit (Macherey-Nagel) while for RNA isolation QIAshredder (Qiagen) followed by RNeasy Mini Kit (Qiagen) was used. Concentration and RNA quality were assessed with NanoDrop 2000 (Thermo Scientific) while for cDNA synthesis PrimeScript™ RT reagent kit (Takara) was used. SYBR® Green (Life Technologies) was used for quantitative real-time polymerase chain reaction (qRT-PCR). All the experiments were assessed on Rotor Gene® 6000 (Qiagen) while results were calculated and normalized to the Hypoxanthine-guanine phosphoribosyltransferase (HPRT) housekeeping gene.

[0094] Partially crosslinked alginate dissolution

[0095] Partially crosslinked alginate dissolution took place at the final differentiation timepoint (Day 21). 0.05M of Ethylenediaminetetraacetic acid (EDTA) (MW:372.24 g / mol, Merk) powder was dissolved in LG-DMEM (Gibco), supplemented with 1% antibiotic-antimycotic (100 units / mL penicillin, 100 mg / mL streptomycin, and 0.25 mg / mL amphotericin B) while the pH was adjusted at 7.4. Differentiated samples on day 21 immersed in the EDTA solution and incubated at 37°C with 5% CO2 and 95% humidity for 60 minutes. Post dissolution the samples were washed 3 times with Tris- HCI buffered saline (TBS). To enhance constructs structural integrity (stand-alone 3d bioprinted scaffold-free constructs) 7 days of additional differentiation was implemented. In Vivo implantation

[0096] Constructs were implanted ectopically in immune compromised mice (Rj :NMRInu / nu) to evaluate their capacity for bone formation upon implantation for 4 and 8 weeks. Mineralization was detected and analyzed with nano-CT. Partially crosslinked alginate constructs without cell was used as negative control and implanted in parallel with the differentiated samples. Upon in vivo implantation the samples were fixed in 4% PFA overnight followed by nano-CT and histological analysis. KU Leuven Animal Research Ethics Committee authorized all protocols related to animal research.

[0097] Nano-CT

[0098] Explant's mineralization compartments were analyzed using nano-CT (Pheonix Nanotom M, GE Measurement, and Control Solutions) with the following parameters: 0.2 mm aluminum filter, diamond target, exposure time 500 ms, skip image 0, images 2400, voltage of 60 kV, current of 170 pA, voxel size of 4.5 pm and mode fast scan with focal spot size 2.42. Image processing and quantification was assessed with CTAn software (Bruker micro-CT, BE) while for the 3D visualization CTvox (Bruker micro- CT, BE) software.

[0099] Histology, Immunostaining and confocal imaging

[0100] Explants were decalcified by immersing in ethylenediaminetetraacetic acid (EDTA) / PBS (pH 7.5) for 14 days at 4 °C after Nano-CT analysis. Paraffin embedding and sectioning occur after decalcification for histological analysis. Histology evaluation was assessed on the sections by staining with Alcian Blue, Masson's Trichrome, and Safranin O. Immuno-histochemistry was performed on PFA-fixed sections of the samples (OSX, IHH, COL1A1, COL2A1, MMP13 and KI67). Antigen epitope retrieval was accomplished using a sodium citrate solution 0.2M for 30 min at 70 °C. Quenching of endogenous peroxidase activity was performed with 3% H2O2 for 10 min. Following, sections were blocked using a blocking solution containing 3% BSA in FBs for lhr and were incubated overnight at 4 °C with the primary antibodies. After overnight incubation with primary antibodies, slides were blocked and incubated with the secondary antibodies for 30 min and peroxidase activity was determined using 3,3'-diaminobenzidine (DAB) (K3468, Dako, USA). Stained histology sections were visualized with a fluorescent Olympus LS microscope. Human size implant biofabrication

[0101] A scale up biofabrication process of the already tested in vitro constructs was performed, to a human dimension implant. Donor derived human periosteum cells (hPDCs) were seeded in collagen microcarriers and undergo proliferation in a GMP grade expansion bioreactor (SCINUS) from 1 million cells to 180 million. During proliferation, integrated sensor technology measures PH, oxygen and pressure at the site of culture. Post proliferation, cells were harvested and mixed with partially crosslinked alginate at 23 million / ml concentration. In total, 2 samples with dimensions height:2cm and diameter:2cm were bioprinted followed by crosslink step at the end of bioprinting presses. A perfusion bioreactor (total volume 250ml) was used to perfuse growth factors towards the differentiation protocol with perfusion rate 23 ml / min. Every 2 days, half of the differentiation media was replaced. Constructs incubated for 21 Days at 37°C with 5% CO2 and 95% humidity. Partially crosslinked alginate was dissolved within the perfusion bioreactor on the day 21. An additional 7 days of differentiation was implemented for 7 more days until the final timepoint (day 28).

[0102] Statistical analysis

[0103] GraphPad Prism 8 (GraphPad Software, Inc., USA) was used for statistical analysis. Data were compared with one-way or two-way ANOVA and Tukey's Multiple Comparison test. P-value statistical difference was set lower than 0.05 (*p < 0.05, **p < 0.01, ***p < 0.001) for all the presenting results.

Claims

Claims1. A method of preparing an implant for bone repair, the method comprising the steps of: a) expanding periosteum derived cells for a period of between 7 and 28 days, b) mixing the expanded periosteum derived cells with a partially cross-linked sodium alginate solution, b) forming the alginate solution and the cells into a shaped construct, c) differentiating the cells in the shaped construct towards chondrogenic differentiation for a period of between 14 and 28 days, d) dissolving crosslinked alginate with a calcium chelating agent.

2. The method according to claim 1, further comprising step e) of cultivating the cells after dissolving the cross-linked alginate for at least 5 days in the medium for chondrogenic differentiation, typically between 5 and 10 days, more typically for 7 days.

3. The method according to claim 1 or 2, further comprising the step of after step b) and prior to step c) further crosslinking the obtained shaped construct with CaCk.

4. The method according to claim 3, wherein the crosslinking is performed by placing the shaped construct in a 1-3 % CaCk (w / v) for between 1 and 5 minutes.

5. The method according to claim 3, wherein the crosslinking is performed by placing the shaped construct in a 2% CaCk (w / v) for 2 minutes.

6. The method according to any one of claims 1 to 6, wherein the partially crosslinked alginate in step b) has a tan(<5) between 0.1 and 0.15, at frequency of 1 Hz or between 0.12-0.13.

7. The method according to any one of claims 1 to 6, wherein between 10 and 40 million cells are added per ml of alginate solution.

8. The method according to any one of claims 1 to 7, wherein between 15 and 30 or between 20 and 25 million cells are added per ml of alginate solution.

9. The method according to any one of claims 1 to 8, wherein the cross-linked alginate in step b) prepared by adding CaCI to a 4% (w / v) alginic acid sodium salt solution until a concentration of between 14 and 25 mM, typically between 17 and 21, more typically 19 mM CaCI is obtained.

10. The method according to any one of claims 1 to 9, wherein forming the alginate and the cells into a shaped construct is obtained by bioprinting.

11. The method according to any one of claims 1 to 11, wherein expanding periosteum derived cells in step a) is performed for a period of between 18 and 24 days, typically for 21 days.

12. The method according to any one of claims 1 to 12, wherein differentiating the cells in step c) is performed for a period of between 18 and 24 days, typically for 21 days.

13. The method according to any one of claims 1 to 12, wherein in the step a) the cells are expanded in a serum comprising basic cell culture medium.

14. The method according to claim 13, wherein the medium comprises 10% foetal bovine serum, antibiotics and antimycotics, the medium being free from growth factors.

15. The method according to any one of claims 1 to 14, wherein in step c) the cells are differentiated in a serum free cell culture medium comprising Bone Morphogenetic Protein 2 (BMP-2), Growth Differentiation Factor 5 (GDF5), Transforming Growth Factor beta 1 (TGF-beta 1), Bone Morphogenetic Protein 6 (BMP-6), and basic fibroblast growth factor 2 (basic FGF-2).

16. The method according to claim 15, further comprising antibiotics, ascorbate-2- phosphate, Proline, Insulin, Transferrin, Selenium, Dexamethasone, and Rho inhibitor, such as Y-27632.

17. The method according to any one of claims 1 to 17, wherein the chelator is EDTA.

18. The method according to any one of claims 1 to 19, wherein the shaped construct is incubated in a medium comprising between 10 and 100 mM EDTA.

19. An implant prepared by the method of any one of claims 1 to 18, for the treatment of a bone defect.

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