Chapter I
The original concept of the current protocol was derived from a scaffold-free cartilage construct made from porcine articular chondrocytes and their high level production of ECM components (Park et al 2006). It was noticed that the scaffold-free cartilage construct could be re-formed into a porous ECM scaffold by freeze-drying process, by virtue of its high water content (~95%) than that of the normal articular cartilage (~85%) (Simon). As shown in the SEM images, a highly porous micro-structure was formed in the cell-derived ECM scaffold, probably comparable to that of PGA scaffold that is used most widely for cell adhesion and growth. Therefore, the cell-derived ECM scaffold might provide a favorable environment for chondrocytes functionally and structurally.
The compatibility of the cell-derived ECM scaffold to chondrocytes was evident from the efficient adhesion and round shape morphology of cells on the ECM scaffold from the early stage. Besides, the rapid increase in the DNA content comparable to that in the PGA scaffold also indicates that the cell-derived ECM scaffold supports well the growth of chondrocytes. The gradual increase in the GAG and collagen contents with time suggests that it provides a favorable environment for the expression of chondrogenic phenotypes. In addition, the synthesized ECMs were uniformly distributed in the cell-derived ECM scaffold, whereas they were observed mainly in the
peripheral area in the PGA scaffold as already shown in many previous studies (Mahnoudifar and Dorn 2005; Mahnoudifar and Dorn 2005; Hu and Athanasiou 2005).
It is speculated that various biological elements resident on the cell-derived ECM scaffolds might have played an important role in this process, such as cytokines, growth factors and potent functional proteins (Badylak 2005). The dynamic interaction between the ECM components and intracellular cytoskeleton is believed to be important in the chondrocyte functions (Hering 1999; Takahashi et al 2003).
One of the most noticeable findings in the study was that the volume of the constructs did not decrease significantly but rather increased at 4 weeks of culture in the cell-derived ECM scaffold. It has long been a greatest challenge to overcome the shrinkage of tissue engineered cartilage. The chondrocytes-seeded type II collagen matrix in vitro was reported to shrink continuously to approximately 50% of its original volume within 4 weeks (Lee et al 2000). In addition, a composite of collagen type I gel and chondrocytes decreased in volume to 50~60% of its initial size after 12 days of culture (Badylak). Natural scaffolds such as fibrin/hyaluronic acid also experienced the same problem of shrinkage in size (Park et al 2005). The contraction of both cell-seeded scaffolds and in vivo implants occurs frequently, which makes the implants unfit for a defect site. Such contraction may even cause the loosening of implants, ultimately leading to separation from the surrounding host tissue. The increase rather than decrease in the construct volume observed in this study was probably caused by more active synthesis of ECM components such as GAG and type II collagen than the
degradation rate of the scaffold. This probably also explain why the compressive strength was gradually increased with culture time in the cell-derived ECM scaffold. In contrast, the compressive strength was gradually decreased with time in PGA scaffold, although the synthesized ECM was increased. Thus, we consider that the physical property of neocartilage is related with the amount and distribution of newly synthesized ECM, and the balance between degradation and synthesis rates of ECM as well. The thickness of the scaffold walls was maintained intact until 2 weeks when it began to degrade, as shown in the histological results. Considering that ECM is degraded by degradation enzymes released by chondrocytes, the cell-derived ECM scaffold can be maintained until the chondrocytes fully expand its number. At this moment they progressively switch their phenotype to synthesis their ECM together with releasing degradation enzyme. Chemical analysis can explain this with showing that GAG synthesis increased rapidly from 2 week of culture on. It is likely that degradation was not dominant until the construct was sufficiently filled with newly synthesized ECM. In contrast, PGA degradation was caused by hydrolysis of water, not by a proteolytic enzyme from chondrocytes. Mechanical study supported this evidence.
ECM construct never decreased its initial mechanical strength, rather increased with time, comparing that PGA construct was getting rapidly weaker regardless of ECM synthesis.
Regarding the size measurements, a new technique was developed using a computer-assisted image analyzing system in our previous report (Choi et al 2006). Compared
with the conventional 2-dimensional (2-D) method, which simply measures the diameter of the constructs (Lee et al 2000; Park et al 2005; Galois et al 2006), the new technique allows volume measurements 3-dimensionally. These results were validated using a hydrostatic weighing apparatus, which produced a margin of error of 0.62~2.2%
(data not shown). Therefore, this method could provide a more accurate measure of the changes in the volume of neoconstructs.
In conclusion, the cell-derived ECM scaffold has a highly porous structure that was suitable for chondrocyte attachment and proliferation. As a cartilage-derived scaffold, its degradation was balanced well by the synthesis of new ECM. In addition, it has the potential to enhance the metabolic activity of chondrocytes as GAG and collagen synthesis. All these favorable influences allowed the construct to not only maintain its size but also produce efficiently cartilage-like structure. These results strongly suggest the importance of the cell-derived ECM scaffold as a novel and promising biomaterial for cartilage tissue engineering.
Chapter II
There have been many reports on the methods of using ECM from the allogenous or xenogenous tissues as a scaffold to produce other organs by tissue engineering (Badylak 2004; Gilbert 2006). However these techniques cannot be applied to the cartilaginous tissue, because the natural cartilage is too compact to provide enough spaces for seeding chondrocytes. In addition, it can cause size and/or shape mismatch, leading to a
step-off in the area of the implantation due to the differences in the surface configurations between the host and donor joints. In order to facilitate the seeding of chondrocytes and maintaining the contents of the cartilaginous ECM, we adopted cell cultivation as a gathering method of ECM. In a previous study, we presented a method for the fabrication of a cell-derived ECM scaffold, and seeding on it of isolated chondrocytes in vitro (Jin et al 2006). Based on this work, we evaluated its feasibility, not only as a cell carrier but also as a structural matrix for cartilage tissue engineering, by implanting the cell-derived ECM scaffold in vivo with chondrocytes. In this study, the cartilaginous tissue was obtained successfully over a short-term period of 3 weeks, by means of a morphological assay, and its biochemical and mechanical properties were evaluated. The experimental specimens retrieved at the 3rd week of post-implantation resembled, in general, normal cartilage.
The most striking feature in this study was that it maintained its volume during cultivation in vivo, even though there was an initial reduction. The shrinkage of tissue-engineered cartilage during cultivation occurs frequently, both in vitro and in vivo.
Theoretically, the reduction of size can make implants unfit for a defect site, and consequently cause loosening of implants and separation from the surrounding host tissue. Lee et al reported that when a cell-seeded type II collagen matrix was used in vitro, it decreased in size continuously until it reached approximately 50% of its original diameter by 4 weeks (Lee et al 200). Galois reported that the collagen type I gel and chondrocyte composites reached 50~60% of their initial size by the 12th day of
culture (Galois et al 2006). A natural scaffold, such as fibrin/hyaluronic acid (HA) also experienced shrinkage with cultivation time (Park et al 2005). Considering that the specimens were very fragile and their sizes significantly decreased at the 3rd week in the control group, it is indicated that the vigorous metabolic activity of the seeded chondrocytes produced ECM sufficiently and the degradation of the ECM scaffold could be balanced by the synthetic rate of the ECM.
Vigorous metabolic activity of chondrocytes can be validated by chemical analysis.
The GAG contents (201.3 µg/mg) in this study was comparable to our previous results using poly glycolic acid (PGA) (201 µg/mg) (Cui et al 2006) or fibrin/HA (189 µg/mg)
(Park et al 2005) as a scaffold at 4 weeks of post-implantation in the nude mouse.
However, the GAG/DNA ratio (60.6) was significantly higher than the values (from 7 to 30) reported previously uising several types of scaffolds such as agarose, alginate, collagen I, fibrin, and PGA cultured in vitro for 40 days (Mouw et al 2005).
According to the RT-PCR and Western blot analyses, type II collagen was markedly enhanced in the experimental group, and it was most prominent at the 3rd week.
However, although type I collagen was also expressed slightly at the gene level, it was not expressed well at the protein level.
It was difficult to clearly distinguish between the newly produced GAG and collagen from the rabbit chondrocytes and those preexisting in the scaffold which originated from the pig chondrocytes. Unfortunately, antibodies specific to the pig ECM proteins are not available without a cross-activity with rabbit ECM. However, considering the
obvious increase of GAG and collagen contents compared with that of scaffold itself and time dependent degradation of scaffold, we speculate that the synthesis of ECM proteins by the seeded rabbit chondrocytes noticeably increased with time in all specimens.
As Badylak reported, cytokines, growth factors, and potent functional proteins were found to reside within the ECM (Badylak 2005). The ECM was recognized quickly as a critically important conduit for the exchange of information among cells. We expected that the ECM scaffold contains as many components as the native cartilage does, because it is produced from chondrocytes isolated from the cartilage itself. Based on the dynamic reciprocity between the ECM and the cells, the seeded chondrocytes must have influenced on the proteoglycan metabolism, as well as collagen synthesis, by maintaining their phenotype and eventually the quality of the tissue-engineered cartilage.
Athanasiou reported that the mechanical strength of the human articular cartilage was 0.53~1.34 Mpa (Athanasiou et al 1991; 1994; 1995; 1998). In the present study, the mechanical strength of the neocartilage tissue constructs increased with implantation time and reached 88% of the normal cartilage strength at the 3rd week of implantation. It was well proportionate to the increment of GAG contents (Fig. 16B). Our laboratory reported that the mechanical strength of tissue-engineered cartilage was 2.25 MPa at 4 weeks in the nude mouse; but rabbit bone marrow MSC and low intensity ultrasound stimulation were included in the study (Cui et al 2006). Thus, it is not stirictly
comparable with the results of this study.
Although Almarza et al reported that the high seeding density of the cells outperformed the low and medium seeding density on ECM synthesis (Almarza and Athanasiou 2005), a large amount of ECM was obtained with our low seeding density at 3.0 x 106/ml. This point is crucial for clinical application, because it is very difficult to obtain a large number of autologous chondrocytes. The DNA content decreased in both groups at the 2nd week post-implantation. This was attributed to the initial adaptation failure of the cells to the environment. Improving the adaptability of cells by in vitro cultivation prior to implantation may produce better results. Since the weight of
the DNA of a chondrocyte is 7.7 ρg, the cell numbers at the 1st week would be approximately 2.9 x 106 in the control group and 3.1 x 106 in the experimental group.
The cells in the experimental group maintained their initial seeded number. Considering that the cells in the control group must have originated from the host tissue, the cells in the experimental group may have contained many host cells initially that were gradually lost by apoptosis or cell death resulted possibly from a shortage of mass transfer. The histological findings that the peripheral region was more active in the accumulation of ECM than the central region may support this idea. Improvements in delivering nutrients and wastes, by such as a dynamic culture in a bioreactor, can provide a promise of a better quality of engineered cartilage.
To the best of our knowledge, this is the first report that demonstrated this comparison by means of volume. In a previous presentation, we validated the accuracy
of our method (Choi et al 2006). When implantation is being considered, a volume measurement of the defect will be an essential step prior to surgery. The volume of the engineered cartilage should meet that of the defect. In this sense, volume measurement is one part of the treatment modality that needs to be developed for clinical usage.
In this study, we demonstrated and evaluated the potential of the cell-derived ECM scaffold as a promising biomaterial structure to produce hyaline-like cartilage in vivo.
The cell-derived ECM scaffold could provide a favorable environment for chondrocytes to maintain their characteristic phenotype and synthesize cartilage ECMs in vivo. We considered that the use of the cell-derived ECM scaffold and chondrocytes formed tissue-engineered cartilage that matured gradually according to the amount of implanted time in the nude mouse, especially after 3 weeks of implantation. Therefore, the cell-derived ECM scaffold is promising for a variety of applications in cartilage tissue engineering including in vivo studies such as an animal model of cartilage defect.
We believe that future studies should be performed on the molecular analysis of the ECM scaffold in order to optimize the culture conditions, such as the addition of growth factors or different culture periods in vitro; and in turn, verify the cultivation in vitro in order to secure implantation within the treated joints. .
Chapter III
In general, a scaffold for cartilage tissue engineering is expected to fulfill prerequisites for structure and biocompatibility, and should be made as a preformed
shape prior to transplantation (Freed et al 1993). For example, the scaffold should be non-cytotoxic and non-immunogenic, guarantee a uniform cell distribution, and maintain cell viability and phenotype. In addition, it should be functionally active to support the synthesis of extracellular matrix components as is required for the development of solid connective tissue (Vacanti et al 1992; Freed et al 1993). Our previous study in vitro and in vivo in nude mouse had proved the cell-derived ECM scaffold could serve as a good scaffold for cartilage tissue engineering by its excellent biocompatibility and supporting well cartilage tissue formation (Jin et al 2006, 2007).
The present study demonstrated two significant points of the ECM scaffold; its feasibility for cartilage tissue engineering in rabbit cartilage defect model and the effect of the maturity of tissue engineered cartilage in vitro on the cartilage regeneration in vivo. The efficiency of the ECM scaffold to support cartilage regeneration of defect was demonstrated by histological observations at 3 months after implantation of engineered cartilage in vitro. The quality of repaired tissues was the best, similar to surrounding native cartilage, when the engineered cartilage was cultured for 4 weeks in vitro before implantation. In addition, the ECM scaffold showed no inflammatory responses in our previous (Jin et al 2007) and present studies, different from the synthetic polymers such as PGA, PLA and PLGA (Bostman and Pihlajamaki 2000; Lu et al 2001).
In many studies on the cartilage tissue engineering, the engineered cartilages have been cultured for various times in vitro before implantation in vivo and no clear consensus has been made about their optimal maturity or culture time in vitro. For
example, chondrocytes were mixed with fibrinogen and implanted immediately to the cartilage defects (Gratz et al 2006) or they were cultured in scaffolds from 2 days to 2 weeks in vitro (Lu et al 2006; Chang et al 2006). We thought that the culture time of engineered cartilages in vitro could determine their maturity such as the ECM expression and mechanical property and affect the quality of cartilage regeneration in vivo. Our results showed that the longer was the culture time in vitro, the better were the quality of engineered cartilage in vitro and cartilage regeneration in vivo. We speculate, however, that optimal culture time should be determined experimentally depending on the scaffold used and cell sources. To our knowledge, this study first reported the effect of maturity of tissue engineered cartilage in vitro on cartilage regeneration in animal defect model.
The fibrocartilage formation by inflammatory cells from subchondral bone occurs during the natural healing process of cartilage defect. In the present study, implantation of the engineered cartilage using the ECM scaffold showed efficient regeneration of cartilage and subchondral bone, and integration of the implants with surrounding tissues.
We speculate that the maturity of the implants might limit the effects of inflammatory cells to cause fibrous tissues. In addition, the physical strength of the mature implants (44% of the native cartilage) resisted the body weight and movement of rabbits, making them not destructed for 3 months.
In the histological observations, the cartilage defect of rabbit without implantation (group 1) showed delayed but spontaneous repair after 3 months at similar levels to
those shown in the implanted groups 2 and 3. The spontaneous repair of cartilage defect might be induced by mesenchymal stem cells (MSCs) and growth factors could be released from the cartilage defect (Chang et al 2006). They might not only induce spontaneous repair of cartilages but also the integration of implants to the surrounding host tissues (Hunziker 1999). We speculate that they could contribute to the regeneration of the cartilage defect particularly in rabbit model, because the thickness of hyaline articular cartilage is approximately 300~400 µm in mature rabbits (Hunziker
1999). Therefore, lager animals such as goat and pig would be better models to evaluate the activity of scaffolds for cartilage tissue engineering.
In conclusion, the engineered cartilage using the cell-derived ECM scaffold could regenerate the cartilage defects, particularly when cultured for 4 weeks in vitro. This result suggests that more mature engineered cartilage can result in better outcome in vivo, probably by providing the better mechanical strength to resist body weight and synthesis of hyaline cartilage-like tissue that can easily integrate host tissues.