## 4 Discussion  
We investigated here the combined influence of a bidirectional perfusion bioreactor and an optimized sequential combination of soluble growth factors for the redifferentiation of HACs in fibrin hydrogel. The device we used in this study, namely the U-CUP bioreactor,  
Figure 6: Scheme of the in vivo study. **(A)**: Following extraction and monolayer expansion, HACs were seeded in fibrin hydrogel and cultured for 3 weeks in a direct perfusion bioreactor. Cartilage constructs were then fixed in fresh human osteochondral blocks with fibrin sealant and hybrid constructs were implanted into the subcutaneous pockets of nude mice for 6 weeks. The stability of the neo-synthesized matrix was assessed by histology. **(B)**: Macroscopic pictures showing (left to right) the cartilage gel fixed in the human osteochondral block from top and side views as well as the hybrid construct implanted under mouse skin. Figure 7: The cartilage construct remains stable after 6 weeks of subcutaneous implantation in nude mice. Following 21 days of culture in a bidirectional perfusion bioreactor in the presence of BIT, fibrin hydrogels seeded with HACs were placed into a cartilage defect made in human osteochondral blocks and then implanted into the subcutaneous pockets of nude mice. **(A)**: hematoxylin and eosin staining of the cartilage construct in the defect after 6 weeks of subcutaneous implantation. **(B,C)**: Immunohistochemistryalyses of the implant. **(B)**: Cartilage implant stained for safranin-O fast green (GAGs) as well as for collagen types I and II. **(C)**: Aggreccan core protein and type VI collagen immunofluorescence staining of the cartilage graft. Proteins of interest were stained with the Alexa Fluor 488-conjugated secondary antibody (red). Nuclei were stained with Hoechst dye (blue) (scale bars: A = 1 mm, B-C = 100 \(\upmu\)m). is a commercially available perfusion bioreactor that takes into account important features of clinical translation, such as the ability to perform maturation of multiple constructs in individual chambers in parallel, automation and ease of use. Here, we applied a relatively low flow velocity based on previous studies. [PERSON] et al. defined an optimal flow rate between 1.6 and 5 \(\mu\)L/s for direct perfusion of bovine chondrocytes seeded in collagen sponge and showed that a higher perfusion velocity (16 \(\mu\)L/s) decreased GAG deposition [37]. Other studies have shown that a flow velocity of 0.13 or 8.3 \(\mu\)L/s increased GAG [15; 16; 18] and aggrecan [16] syntheses in agarose hydrogels by animal chondrocytes compared to static cultures. However, [PERSON] et al. reported that a perfusion velocity of 1.7 \(\mu\)L/s reduced the production of GAGs and link protein by human chondrocytes in alginate hydrogels [17]. Here, our results demonstrate an increased production and improved quality of neo-synthesized cartilage matrix by HACs in fibrin hydrogels with low perfusion velocity (2.5 \(\mu\)L/s) compared to static cultures. Perfusion preconditioning was applied to HAC-seeded gels of clinically relevant diameter, since 10 mm is the minimum defect diameter recommended for ACI [6; 7] and is often reported as the average cartilage defect diameter in clinical studies involving ACI [38]. Perfusion clearly increased the quantity of cartilage ECM deposited in the large engineered tissues, with up to three quarters of the graft cross-section positively staining for GAG deposition. However, a delayed maturation of the core was indicated by ectopic cartilage ECM-positive areas, whereas the periphery of the gel seemed to have achieved its maturation as evidenced by homogeneous cartilage ECM deposition. This may be explained by an insufficient mass transfer of nutrients and growth factors to the core, as well as by lower fluid-induced shear stresses within the scaffolds which could have understimulated the cells. In this study, we paid particular attention to the nature of the newly synthesized matrix.