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Chondrocytes in articular cartilage synthesize collagen type II and large sulfated proteoglycans, whereas the same cells cultured in monolayer (2D) dedifferentiate into fibroblastic cells and express collagen type I and small proteoglycans. On the other hand, a pellet culture system was developed as a method for preventing the phenotypic modulation of chondrocytes and promoting the redifferentiation of dedifferentiated ones. Because the pellet culture system forms only one cell aggregate each tube by a centrifugator, the pellet could not be applied to produce a tissue-engineered cartilage. Therefore, we tried to form chondrocyte aggregates by a rotational culture, expecting to form a large number of aggregates at once. In order to increase cell–cell interactions and decrease chondrocyte–material interaction, dishes with low retention of protein adsorption and cell adhesiveness were used. In addition, rotational shaking of the dish including cells was attempted to increase the cell–cell interaction. The shaking speed was set at 80 rpm, so the cells would be distributed in the center of the dish to augment the frequency of cell–cell contact. Under these conditions, bovine articular chondrocytes started aggregating in a few hours. At 24–36 h of rotational culture, aggregates with smooth surfaces were observed. Parameters such as increase of culture time and addition of TGF-β controlled diameters of the aggregates. There were many fusiform cells at the periphery of the aggregates, where the cells tended to form a multilayered zone in cross sections. In addition, lacune-like structure, which was almost the same as pellet culture, was observed. It was found that the internal structure of the aggregates was similar to that of pellets reported previously. Therefore, the aggregates formed by a rotational culture could become an essential component to make tissue-engineered artificial cartilage.

Although rapid formation of a smooth inner surface is important in constructing an artificial vascular graft, a conventional model that uses a biodegradable polymer such as polyglycolic acid needs long-term culture to form it. In another model, which uses collagen gel, it is reported that prompt formation of the smooth inner surface was achieved. But the mechanical properties were not suitable, resulting in rupture under high pressure at the arterial level. Therefore, we propose a new artificial vascular graft model made of biodegradable polymer, gel, and cells. At first we manufactured an artificial vascular graft (i.d. 5 mm, o.d.7 mm) consisting of poly-l-lactic acid (PLLA) with open pore structures by using gas-forming methods. After mixing human normal aortic smooth muscle cells (SMCs) with type I collagen solution, pores of the PLLA scaffold were filled with the mixture. The collagen mixture was made into gel in the pores of the PLLA scaffold, incubating at 37°C. WET-SEM analysis showed that the prompt formation of a smooth inner surface was achieved in the new model. The ratio of incorporation of SMCs into the artificial vascular graft became approximately 100% by using the cell–collagen mixture, whereas only 40% of SMCs were trapped in the conventional model where SMCs were inoculated as a cell–medium suspension. Therefore, it was suggested that the new artificial vascular graft model was superior in smooth inner surface formation and cell inoculation, compared with conventional models using either biodegradable polymer or gel.

Tissue engineering approaches have been clinically tried to repair damaged articular cartilages. It is an essential step to uniformly seed chondrocytes into 3D scaffolds in order to reconstruct tissue-engineered cartilages in vitro, but the tissue engineering could not have been provided with efficient cell seeding methods. Type I collagen is clinically used and known as a cytocompatible material, having recognition sites for integrins. Collagen gel encapsulating chondrocytes has been tried for making regenerated cartilages, but it is found difficult to have the gel keep its original shape after long-term culture, because of shrinking. On the other hand, 3D scaffolds, either of a nonwoven structure or a sponge-like structure, involve difficulty in that chondrocytes could not be uniformly seeded, although they have adequate initial mechanical properties. In this study, by combining collagen gelation with a nonwoven PLLA scaffold, we achieved uniform cell seeding into the 3D scaffold. Bovine articular chondrocytes were mixed with type I collagen solution, and the solution was poured into the nonwoven PLLA scaffold (1.5 mm thick, f 15 mm). The collagen–chondrocyte mixture was made into gel at 37°C for 1 h. The 0.39% collagen mixture was viscous enough to prevent cells from precipitating during gelation. Almost all chondrocytes were able to be incorporated into the PLLA scaffolds by mixing with collagen solution and subsequently making into gel, while 30–40% of the chondrocytes seeded as a cell suspension were not trapped into the PLLA scaffolds. The method presented, where chondrocytes were mixed with collagen solution, and the mixture was incorporated into a 3D scaffold, then made into gel in the scaffold, could serve as an alternative for in vitro cartilage regeneration, also simultaneously having the advantages of both materials.