Jacobs Journal of Bone Marrow and Stem Cell Research

Emerging Concepts for Articular Cartilage Regeneration

*Luminita Labusca
Department Of Orthopedic Surgery, County Emergency Hospital, University Of Copenhagen, IASI, Romania

*Corresponding Author:
Luminita Labusca
Department Of Orthopedic Surgery, County Emergency Hospital, University Of Copenhagen, IASI, Romania

Published on: 2018-06-06


Articular cartilage diseases still represent an unmet clinical need. After decades of basic, translational and clinical research, cartilage engineering still struggles in finding the best modalities for fabricating functional grafts that will endure the test of time. The classical Top Bottom approach in tissue engineering (TE) consisting in putting together cells biomaterials and growth factor is challenged by the emergence of novel ‘bottom top “concepts and technologies. Bioprinting and scaffold free (SF) cartilage engineering even though technically and conceptually different, take advantage of behavior of biological systems remembering of developmental stages. By mimicking the developmental niche it is sought that hierarchic organized and functionally adapted grafts for cartilage engineering could be produced. Several clinical trials using SF fabrication are already going on opening the perspective bottom top approaches could deliver clinically relevant solution for joint surface regeneration in the near future.


Articular Cartilage; Bioprinting; Scaffold Free; Self Aggregation; Self Assembly

Copyright: © 2016 Luminita Labusca


The field of articular cartilage repair attracts considerable interest from the part of basic scientists, bioengineers and medics. Numerous teams are engaged in the quest for finding the most viable and enduring method for restoring a damaged joint surface. However, after several decades of intense research, cartilage diseases still represent an unmet clinical need. Several explanations for the limited tendency to heal of articular cartilage are the functionally adapted structure as well as the inability to initiate inflammation-based repair processes in the case of injury. As an ultra-specialized tissue adapted to function under important mechanical stress, cartilage is aneural, avascular and a lymphatic with low cellularity and high content of extracellular matrix (ECM) [1]. Cartilage destruction is seldom life-threatening, however, it produces important suffering and disability imposing a consistent burden on healthcare systems and increasing health-related expenditures [2]. Regenerative medicine (RM) introduced the concept of complete structural and functional restoration of tissues organs and even systems that could be possible by employing breakthrough discoveries in the field of molecular cell and developmental biology as well as cutting edge technologies such as Nanomaterial fabrication, bioprinting, in vivo cell tracking and imaging [3]. It is sought that using RM specific tools complete structural regeneration and functional rehabilitation of affected joint would be possible in the near future. From this perspective, the non-complicated structure of a tissue composed only by ECM fibbers and one cell type (articular chondrocytes) appeared relatively simple to be reproduced. 

However, cartilage engineering has proven to be a challenge. While many research groups all over the world are engaged in basic, translational or clinical research in this field, viable solution for multipurpose, time enduring replacement of articular surface are still awaited. The major characteristic of articular cartilage, the remarkable biomechanical endurance has proven not trivial to reproduce by engineered counterparts. Despite the apparent structural simplicity, cartilage has a distinct multi-layered distribution of cells and ECM fibers. Zonal particularities of both cell shape and density together with the topographical arrangement of collagen fibers and proteoglycan aggregates are factors that determine tissue resilience, shear resistance or compressive stiffness [4]. It is expected that a more accurately mimicking tissue heterogeneity, fibber and cell distribution would lead to producing cartilage grafts of superior functionality. Adding to the complexity of the problem, the stable integration of the engineered graft within the surrounding tissue as well as graft attachment the underlying subchondral bone, are issues still to be resolved in the case of biosynthetic grafts. Recently, the classical paradigm of using cells, biomaterials and bioactive molecules [5] mixed together to produce implantable tissues is shifting towards new engineering concepts that are based on advanced technology as well as on the deepened understanding of developmental biology. Without being exhaustive, this review will introduce 3D bio-printing and scaffold-free approaches as novel concepts and technologies for engineering articular cartilage.


Cartilage bio printing - technology leads the way

For the last decades, various combinations of cell, scaffolds and bioactive molecules are used aiming to produce functional cartilage grafts. Launched in the early 90 ties, scaffold-based top-bottom approach is still the most popular modality to engineering tissues [6] Choices regarding cell source i.e. adult or stem cells, composition, rate of biodegradability and methods of scaffold manufacturing, use of bioactive molecules - growth factors or cytokines- as well as the method of delivery are largely variable (for review see Ref [7]). Most of these strategies involve using a piece/part of a supportive structure (scaffold) manufactured as a non-living compound that will receive, at some point of the engineering process cell population(s), bioactive molecules or a combination of both. Bioprinting technology is shifting the concept of graft fabrication towards the production of a structure that includes inert as well as living compounds deposited together in one single step of the manufacturing process [8]. Bioprinting, regarded as a form of “additive manufacturing”, requires layer by layer deposition of a mixture of cell and biomaterials, therefore, composing objects from the bottom to the top [5]. In this process a three-dimensional (3D) bioprinter is a computer controlled robotic system that creates 3D objects based on a “blueprint”, a computerized model of the structure to be produced [9]. The process uses “bio-ink”, that is a modular building block composed of cells and viscous biomaterials deposited together in a layer by layer manner to construct the desired tissue or organ shape. The printed structure has to undergo maturation in special designed bioreactors [10]. The deposited cells fuse together in a process similar to the tissue fusion occurring during embryonic development (see as well biomimicry chapter).

One of the main advantages of 3D printing is the reduced time required for obtaining and modifying product characteristics allowing for “rapid prototyping” and increased versatility. Moreover, geometrically complex shapes, layers and topographical arrangements of compounds can be achieved potentially allowing for precise reproduction of the native tissue structure and graft customization. It is expected that with bioprinting, cartilage zonal cell and fiber distribution could be reproduced, therefore improving graft mechanical properties. The biosynthetic tissue could be manufactured to fit with high precision the cartilage defect or surface to be grafted. Digital Imagistics such as computed tomography (CT) or nuclear magnetic resonance (NMR) could be used to derive the ‘blueprint’ for fabricating customized grafts. Among printing technologies that have been demonstrated to preserve cell viability and proliferative capabilities are fused deposition modelling [11] thermal inkjet printing [12] selective laser sintering [13] stereo-lithography [14] extrusion printing gels [15] Hydrogels in combination with either adult (chondrocytes) or progenitor cells are the most commonly used starting materials while inkjet and pneumatic extrusion the most commonly used technologies for cartilage bioprinting [16]

Hydrogels appear to be the suitable material for bioprinting cartilage due to their ability to undergoing a phase change from liquid to solid by cross-linking Hydrogels are water swelling - water-insoluble cross-linked networks that provide the 3D environment and are able to maintain the high water content favorable for cell survival and proliferation. Natural (such as collagen, chitosan, alginate, hyaluronic acid (HA) silk proteins as well as synthetic polymers such as Poly(vinyl alcohol) PVA, Poly(ethylene glycol) (PEG) and combinations have been widely used as hydrogels for cartilage engineering [17]. These materials and their derivatives were shown to exhibit swelling and lubricating properties similar to the native tissue and to promote the chondrogenic phenotype for the encapsulated cells [18]. In the presence of adequate initiator molecules, phase shifting by cross-linking can be induced chemically, thermally or by exposure to light or UV [5]. Various combinations of hydrogels, synthetic biomaterials and cell populations with or without growth factors have already been reported to generate stable cartilaginous tissue by bioprinting technologies. Additive manufacturing with a multi-head deposition system (MHDS) was used to fabricate 3D cell-printed scaffolds using layer-by-layer (LBL) deposition of polycaprolactone (PCL) containing transforming growth factor-β (TGFβ) and alginate encapsulated chondrocytes. When implanted in the dorsal subcutaneous space of female nude mice the bio-printed tissue was shown to express collagen type I characteristic for articular cartilage [19]. In another report, anatomically shaped fibrous and elastic cartilage structures, (a human ear and sheep menisci) were 3D printed using MRI and CT images as blueprints. Bio-ink formed of human chondrocytes encapsulated in nano cellulose-alginate allowed for good cell survival at 7 days after printing [20]. 

In the quest for bio-inks that have good rheological properties as well as increased cytocompatibility and product versatility, thermoresponsive polymer poly (N-isopropylacrylamide) grafted hyaluronan (HA-pNIPAAM) was blended with methacrylate hyaluronan (HAMA). The HA-pNIPAAM compound provided fast gelation and post-printing structural fidelity, ensuring long-term mechanical stability upon photocrosslinking. Bovine chondrocytes displayed good viability after elution of HA-pNIPAAM from the scaffold proving that this method could be used for printing of stratified cartilage constructs [21] .

The combination of bio-printing with other advanced manufacturing technologies was found to improve cell viability within the construct as well as the biomechanical properties of the deposited tissue.

An innovative approach in this respect used a combination of bioprinting and microcarrier technology. Mesenchymal stromal cell (MSC) in polylactic acid microcarriers, encapsulated in gelatine methacrylamide-gellan gum bio-inks were used to obtain a bilayer construct reproducing the cartilage layer as well as the subchondral bone. Increased cell concentration and viability could be obtained by bioprinting while microcarrier encapsulation improved the mechanical characteristic of the hydrogel layer, facilitated cell adhesion and MSCs osteogenic differentiation in the subchondral bone layer [22].

A hybrid system combining electrospinning of polycaprolactone fibbers with inkjet printing of rabbit elastic chondrocytes suspended in the fibrin-collagen hydrogel was used to fabricate a 1 mm thick, 5 layered cartilage like construct. Good cell viability and proliferation as well as in vitro and in vivo cartilage matrix deposition and enhanced mechanical properties compared to printed alginate or fibrin–collagen gels alone was reported [23].

Due to their superior proliferation and differentiation capability, the use of stem cells in the bioprinting process is appealing. Using the appropriate technology, the same stem cell source could be employed to fabricate heterogeneous tissues such as osteochondral grafts.

Inkjet bioprinting of poly(ethylene glycol) dimethacrylate, gelatin methacrylate, and human MSCs were used to obtain grafts of increased mechanical resistance. Layer by layer deposition of simultaneously photocrosslinked PEG-GelMA hydrogel scaffolds resulted in the formation of constructs with mechanical properties similar to native tissue. Moreover, the scaffold was shown to promote osteogenic and to have positive effects on chondrogenic differentiation of hMSCs of potential interest in producing osteochondral grafts [24]. Addressing several issues with the manufacturing processes such as low cell survival and frequent clogging of the print head, the same team further developed the procedure combining multiple steps of scaffold synthesis and cell encapsulation successfully combined into one single step procedure allowing for minimal exposure of cells to UV during photopolymerization. hMSCs osteogenic and chondrogenic differentiation was improved while the bio-printed bone and cartilage were shown to express mineral and cartilage matrix deposition respectively, as well as increased mechanical properties. Moreover, the bio-printed PEG-peptide scaffold could inhibit hMSC hypertrophy avoiding cartilage matrix ossification, a relatively common drawback that hampers MSC derived chondrocytes to preserve the specific phenotype [25].

Cell contact with natural ECM is a crucial factor in maintaining cell viability and phenotype; Natural dECM supports progenitor cell survival and differentiation and this was found to be true for obtaining printed constructs; cell encapsulation within their specific decellularized ECM (ECM) hydrogel prior to the deposition facilitates formation of tissues of increased viability and functionality [26]. Bioprinting using dECM bio-ink was tested in order to obtain a versatile microenvironment that could better support the growth of three-dimensional bio fabricated structures. Thus, 3D open porous structures of MSCs cell-laden dECMs with polycaprolactone (PCL) framework were printed using a multi-head tissue/organ building system. This allowed the formation of tissue-like structures with good cell viability, specific lineage differentiation and increased de novo ECM production. Moreover, with this approach, different ECM types could be deposited, allowing for composing heterogeneous structures of importance for printing complex tissue architectures [27].

Laser printing (Laser-induced forward transfer –LIFT- ) was used to generate MSCs based scaffold-free grafts. Predifferentiated MSCs were shown to survive the printing process and were deposited in high density that allowed for chondrogenic differentiation. Authors conclude that using this method autologous complex scaffold-free bone and cartilage grafts could be developed [28].

Maybe one of the most exciting development is the perspective of in situ printing. Direct cartilage repair with engineered tissue that could closely reproduce native cartilage properties to the site of the lesion without damage to the healthy tissue but allowing for improved graft integration, is very appealing. Using a set up with simultaneous photopolymerization, Poly (ethylene glycol) dimethacrylate (PEGDMA) and human chondrocytes were printed to repair defects directly into osteochondral plugs (3D bio paper). The printed tissue was shown to attach firmly with surrounding tissue displaying enhanced proteoglycan deposition at the interface of implant and native cartilage in Safranin-O staining as well as enhanced interface failure strength during push-out testing [29].

Cartilage bio-printing faces a lot of challenge some of them similar to other organ printing technologies, some of them related to the particularity of the tissue. Common challenges related to the “maturity” of the technology include the necessity for an adequate computer-assisted designed (CAD) organ blueprint and in silico modelling of the tissue self-assembly, step by step design of bio fabrication process, development of bioreactors that allow for non-invasive monitoring of post-printing maturation (for review see [30]). Logistical challenges are related to the imperative of developing complex multidisciplinary teams that include experts in the field of biomedical engineering, biologists and computer scientist, access to wide spectrum of ‘raw materials’- cell sources and biomaterials as well as to bio-fabrication equipment and the availability of a large panel of destructive or/and nondestructive product characterization tools. The perspective of in situ bio-printing, in particular, has to overcome issues related to the bulkiness of the equipment that might prove difficult to install in an operating room. Bio-printing cartilage faces specific issues mainly related to the availability of the cell source and the long term stability of the graft. An adult cell can be obtained in limited number from cartilage biopsy, however, the in vitro expansion potential, as well as viability after printing, remains limited. The use of MSCs of autologous or allogeneic source with superior proliferative and differentiation is appealing, however, it is currently still restricted by the cumbersome process of ethical and regulatory approval required for in translating stem cell-based procedures from laboratory to clinical applications. Bio-printed cartilage graft integration within the host as well as mechanical stability that accounts for long term endurance and functionality of the engineered tissue within the context of the joint complex have to be further investigated. Currently, there are no data to support the superiority of printing technology in providing cartilage repair solutions. Regardless of the current limitations, the technology is developing at high speed and it is expected that medium and long term in vivo studies and potentially first clinical trials will become possible in the near future. 

Scaffold free cartilage engineering – the biological approach–

In the same line of challenging the classical triad for engineering tissues and organs, scaffold-free approach (SF) is emerging as a modality of fabricating tissue grafts without the use of a supportive biomaterial. Taking advantage of cell natural ability to bond together, to express ECM proteins and to respond to stimuli, SF uses different methods of agglomerating cell populations inducing them to produce their own ECM in order to compose an implantable graft. Several drawbacks in using biomaterials as scaffolds have generated the SF approach. The presence of a scaffold can alter the phenotype of the cell in contact with it, the degradation process seldom synchronizes with de novo tissue deposition and can generate by-products that cannot easily be cleared apart. Its bulkiness can as well interfere with cell mechanosensitisation by physically creating a stress shield [31] (Table 1). SF also denominated scaffold -less tissue engineering is defined as “any platform that does not require cell seeding or adherence within an exogenous, three-dimensional material’’ and has already been used for engineering articular cartilage, meniscal tissue, intervertebral discs [32]. A large number of cells are put together in order to generate a situation that is similar to the processes of condensation and differentiation that takes place during cartilage development. Cells are expected to produce ECM composing de novo cartilage that will expand and maturate under the effect of various external stimuli such as growth factors, enzymes, mechanical stimulation. SF technology encompasses two distinct modalities of cell agglomeration: self-organization and self-assembly, the main distinction between the two is the presence or absence of external forces imposed to the system (for review see Ref [32]). A common trait is that both techniques require rather important starting number of cells, therefore the necessity of a relevant cell source when aiming productization and clinical translativity is crucial [33]. Recently, SFTE of various tissues in vitro has been achieved using new techniques of induced microgravity on ground-based facilities. Microgravity obtained using devices such as the 3D Random Positioning Machine (RPM), the 2D FastRotating Clinostat (FRC) or the Rotating Wall Vessel (RW) was demonstrated to enhance 3D structure formation as well as the size of organ-like aggregates [34]. Human umbilical CD34 positive cord blood cells or the endothelial cell line EA.hy92 were proven to display increased growth and differentiation into a vascular phenotype as well as increased expression of genes involved in signal transduction, angiogenesis, cell adhesion, membrane transport or serine synthesis as assessed by genome-wide microarray analysis when cultured in RPM microgravity conditions as compared to static cultured cells [35].

Three distinct directions for SF cartilage engineering is evolving: cell sheet and cell aggregation as pertaining to the self-organization category and the self-assembly based engineering.

Cell sheet technology

Cell sheet and cell aggregation commonly use a form of external force or external manipulation coaxing living elements to produce compact tissue-like structures. For

Table 1. Clinical trials investigating Scaffold free methods for cartilage engineering (www. clinical trials.gov accessed February 2016)

sheet formation, cells are expanded in vitro at high density in monolayer for extended periods of time allowing them to form a continuous layer. The structure can be lifted as a whole either thermally or by enzymatic digestion and further manipulated (rolled, dropped or molded) to undergo tissue fusion. Sheets subsequently fuse together in a process similar to tissue fusion during embryologic development by means of cell-cell and matrix to matrix contact and ECM remodeling [36]. Cell sheet technique for producing implantable cartilage grafts has been successfully tested in animal models of cartilage defects. Chondrocyte sheets prepared in a temperature responsible culture dish were used to repair osteochondral defects in minipigs with good results respective to neocartilage formation and tissue integration enabling the development of the procedure for human chondrocytes and potential clinical applications [37]. Adipose-derived stem cells (ASCs) are an appealing source as they can provide the high number of cellular elements required for the aggregation process. However, the extended in vitro culture required to drive ASC chondrogenic differentiation within the forming cell sheet can be a problem when translating to clinical applications. Incorporation of transforming growth factor-β1 (TGF-β1)- loaded gelatin microspheres was reported to result in sheets with similar chondrogenic properties (GAG production) to the ones treated with exogenous TGF-β1. This fact was reported as a proof of concept for speeding the formation of an implantable graft for clinical applications [38].

Indeed, cell sheet technology is already translating to clinics. RevaFlex, a product based on expanded allogenic juvenile chondrocyte sheets is currently undergoing a Phase III clinical trial after reporting good results from a Phase I/II study in 6 out of 9 patients treated as assessed by MRI and second-look arthroscopy [39]. A clinical trial using autologous human chondrocytes layered as cells sheets in a two-step procedure (arthroscopic cartilage biopsy harvest and open cell sheet implantation) is reportedly going on in Japan for the treatment of full-thickness cartilage defects of the knee [36].

One of the main inconvenient with the procedure is that it requires extensive cell manipulation that can prove difficult

Table 2. Advantages and disadvantages of top-bottom and bottom -top approaches in tissue engineering (TE)

to standardize and scale up. As it is the case with other tissue engineering methods, problems related to nutrient diffusion for the fabrication of thicker tissues needs to be resolved when producing larger cartilage grafts. The use of adult chondrocytes to form cell sheets is limited by cell dedifferentiation while in monolayer and the shift towards a fibroblastic phenotype that has deleterious consequences on the quality of the engineered tissue. Silicon rubber based continuously expanding culture surfaces [40] or employing mesenchymal stem cells of various origins as cell source has proven efficient in avoiding the problem of chondrocyte dedifferentiation and proved successful in generating good quality implantable grafts [36].

As with other cartilage engineering technologies, graft integration within the host tissue and to the subchondral bone remains a problem. A possible solution would be the attachment of the cell sheet generated cartilage constructs to support capable to mediate osseous integration of the cartilage graft. Scaffold-free chondrocyte sheets were successfully integrated to porous titanium demonstrating neocartilage tissue ingrowth within the pores, a method with potential in providing a graft with good fusion with underlying bone [41].

Cell aggregation is a commonly used laboratory technique that employs rotational force in order to obtain cell conglomerates. The technique is variable from using high speed (400-500g) centrifugation for short amount of time (3-5 minutes) that results in a pellet culture to low speed (tens of revolutions per minute) and high duration (14-21 days) used in a laboratory or industrial rotational bioreactors. Aggregation is used to agglomerate cells for inducing chondrogenic phenotype in progenitor cells [42] to re-differentiate chondrocytes dedifferentiated after long time two-dimensional expansion [41] or as an SF method to engineer cartilage. Cartilage grafts obtained by cellular aggregation increase steadily in size with culture time so that the growth of a tissue starting from a given shape might be mathematically predictable. Moreover, mechanical properties could be adjusted by mechanical stimulation during tissue growth, therefore, enhancing the functional quality of the graft [43]. As a plus, the technology is amendable for screening various soluble factors on stem cells of different origins in a high throughput manner suitable for laboratory as well for industrial applications [44]. MSCs appear to be the most appealing cell source for engineering SF cartilage by aggregation. MSCs (pellet culture) exposed to TGF β1 were induced to agglomerate into condensed mesenchymal cellular bodies (CMBs) in order to undergo chondrogenic differentiation. Over 5 weeks, the procedure generated cartilaginous tissue in a manner considered to mimic mesenchymal condensation leading into chondrogenesis during embryonic development. The cartilage of anatomic shape and of clinically relevant dimensions could be produced by this method in the extent of 5 weeks. Engineered grafts were found to having unprecedently comparable mechanical properties to native tissue (Young's modulus of >800 kPa and equilibrium friction coefficient of <0.3) [45].

SF cell aggregation technology is already translating to clinical applications. A phase III clinical study is currently going on investigating an autologous chondrocyte aggregates/spheroids based commercial product (Chondrosphere) for the treatment of cartilage lesions of the femoral condyle. A previous PhaseI/ II clinical trial using Chondrophere for the treatment of International Cartilage Repair Society (ICRS) grade 3 or 4 cartilage defects resulted in significant improvement of Lysholm, International Knee Documentation Committee [IKDC], SF-36, Tegner scores as well as in advanced defect filling as assessed by MRI [46].

From a cartilage engineering perspective, cell aggregation can be used as a step during other engineering approaches or as a form of SF fabrication. Cell aggregation drives chondrogenic differentiation of mesenchymal stem cells or re-differentiation of dedifferentiated chondrocytes. Cells can be further dissociated and used to form cartilage by seeding them on a scaffold in a classical top-bottom approach or in bioprinting [47] Cell aggregates can be used to directly to fill cartilage defects as a stand-alone SF cartilage TE method [48].

Problems with cell viability in the initial period of cell suspension, reduced cell proliferation, and viability in larger aggregates are several issues of the technology. Formation of smaller aggregates to be used as building blocks for with other engineering methods as well as the direct use of cell aggregates to fill a cartilage defect partially addresses these issues [49].

Self-assembly is a method of forming tissue-like structures without the use of external forces.

Cell self-organization occurs following the thermodynamic principle of minimization of free energy by means of cell-cell interaction. Typically when (normal adherent) cells are seeded on nonadherent substrates a sequence of processes that recapitulate tissue formation takes place forming constructs with functional properties relevant to the native structure [31]. Thus, a high-density suspension seeded in a non-adherent mold will lead to cell coalescence, ECM formation followed by ECM maturation, therefore, composing functional tissues. The method is considered to be highly biomimetic as it practically recapitulates the stages of tissue formation during embryonic development. Using molds of appropriate structures and sizes, engineered tissues can be fabricated to meet the clinical needs.

In the field of cartilage engineering, SF fabrication by self-assembly is a method that has attracted interest in recent years. As a proof of concept, self-assembly of bovine chondrocyte within agarose molds resulted in constructs displaying biochemical, histological and mechanical properties similar to hyaline calf articular cartilage [50].

Stem cells are of increasing interest in self-assembly technology for obtaining cartilage grafts. Infrapatellar fat pad-derived stem cells on polyethylene terephthalate (PET) transwell membranes transiently (first 21 days of culture) supplemented with TGF-β3 lead to the development of constructs displaying similar GAGs content with those obtained by agarose gel encapsulation. However, when normalized to tissue weight, self-assembled constructs displayed higher ECM accumulation per unit [51].

Modulation of culture condition has been shown to improve the parameters of engineered cartilage. Human bone marrow MSCs submitted to chondrogenic differentiation in monolayer in media supplemented with growth differentiation factor 5 (GDF-5) underwent a self-assembly process under hypoxic conditions, the method resulted in the formation of cartilage-like structures with significant higher ECM content compared to normoxic conditions and having higher structural similarity with the natural tissue in terms of fibbers disposition [52]. Polydimethylsiloxane (PDMS)-based culture chip having arrays of microcavities were shown to generate large numbers of uniform spheroid like structures with narrow size distribution and to enhancing chondrogenic differentiation potential in ATDC5 cells, a modality of generating 3D cellular constructs with high functionality for tissue engineering [53].

Tissue-like structures that recapitulate anisotropic matrix organization could be obtained by self-assembly of a mixed cell population and geometrically shaped molds. 50:50 bovine articular cartilage and meniscus cells were grown in temporomandibular (TMJ) shaped molds under biomechanical (passive axial compression) and bioactive molecule stimulation (chondroitinase-ABC and TGF-β1) resulted in a graft with collagen fibril alignment, Young’s modulus, and ultimate tensile strength similar to the native tissue [54].

The principal inconvenient with the self-assembling process is the high starting number required for the procedure within tens of millions of cells/ mL. Potentially reduced cell viability in the initial stages of assembling process that expose cells to minimal cell ECM interaction, as well as issues with nutrient diffusion within larger constructs, are other problems to be resolved by this technology [55]. However, despite the fact that higher amount of cells were required, a comparison between self-assembled (SA) and agarose hydrogel encapsulated self-aggregated chondrocytes concluded that SA constructs generated regularly formed shapes with higher ECM content in less time. The faster generation of a more hyaline-like tissue could counteract the necessity for a rich cell source when designing strategies for clinical applications.




Common denominator of bottom top approaches – biomimicry

Engineering 3D tissues starting from cellular components could be considered as bottom top approaches. Scaffold-free technologies, as well as bioprinting, are based on several properties of individual elements that lead to the formation of more complex structural and functional systems. Some of these properties determine cell behavior highly similar to the ones occurring developmental processes of tissue and organ formation. Tissue fusion represents the process by which isolated cell population come in contact and adhere by means of cell-cell interaction, cell-matrix interaction and ECM remodeling. During developmental stages, tissue fusion occurs in neural tube formation, cardiac development or skeletal patterning [56]. Bio-printing is a modality of directed self-assembly where the separate elements deposited are allowed to fuse together during the maturation process. Scaffold-free technologies use either an external force -in the case of self aggregation-either the principle of free energy minimization-in self-assembly- [57] to induce close cellular proximity that allows for tissue fusion. Engineering cartilage or any other tissue using bottom top technologies ultimately depends on a process of tissue fusion. By any of these methods, separated cells come in contact and fuse together forming continuous grafts. By closely reproducing the developmental process of cartilage formation during mesenchymal condensation, it is expected that bottom top engineered grafts would display a level of cell heterogenicity and ECM proteins organization that mimick natural tissue [58]. In this context, two important elements are needed in order to modulate the process of cartilage formation and to mimick the developmental niche: signaling molecules and mechanical stimulation. Growth factors of TGF β familly, bone morphogenetic protein (BMP) or insulin growth factor (IGF) have been shown to influence chondrogenic differentiation of progenitor cells, matrix production as well as neotissue mechanical properties (for review see [59]). Enzymes such as Chondroitinase ABC and Lysyl oxidase (LOX) were proven effective in enhancing neotissue mechanical properties by reproducing native tissue remodeling during growth stimulating as well as developmental pathways involved in tissue maturation [32].

During embryonic development, cartilage formation and maturation depends on the presence of appropriate mechanical stimuli. Mechanical forces such as compression, tension shear or combinations are widely applied for both top bottom and bottom top cartilage engineering. Mechanical stimulation of engineered tissue aims to reproduce the developmental environment enabling the production of functional grafts by influencing the chondrogenic phenotype, matrix production and tissue maturation [60].

A place among cartilage repair solutions – are bottom top technologies clinical ready?

The current surgical option for focal cartilage defects include microfracture, osteochondral autograft, and allograft transfer, autologous chondrocyte implantation (ACI) or matrix-assisted autologous chondrocyte implantation with documented clinical results Such techniques have become increasingly popular among orthopedic surgeons and harvest good results in selected clinical situations [61,62] However none of the mentioned techniques has the capability of restoring the structural and functional integrity of a joint surface, therefore, the clinical results are prone to deteriorate in time. Cartilage bio-printing and scaffold-free approaches attempt to reproduce the mesenchymal condensation phase during cartilage development. This enables a more reliable reproduction of gene expression dynamics closely followed by increased mechanical properties of printed [16] or scaffold-free generated tissue. Moreover, simultaneous cartilage-subchondral bone regeneration could eliminate issues related to cartilage-bone interface and could enable restoration of larger joint areas [63]. Bio-printed cartilage is not currently clinically tested; however, this approach is expected to become the technology of the future at least for large joints such as knee or hip [64]. Important improvements regarding equipment and workflow are needed before bioprinting develops into a clinical ready technology. The choice of most suitable cell-hydrogel combinations will need to take into account cell availability, the modified phenotype in relation with bio-ink as well as immediate and long term properties of the printed tissue. Results from currently going on clinical studies involving scaffold-free approaches are awaited in order to generate the evidence recommending them as therapeutic options among existent cartilage repair procedures. 

Bottom top technologies – future directions

Of increasing importance among cartilage engineering procedures, bioprinting, as well as SF technologies, aim to produce cartilage grafts that closely resemble architectural organization and functional properties of native tissue. While they are similar in manipulating fundamental principles of cell behavior and developmental biology, bottom top approaches differ in their modality of resolving the main issues encountered in cartilage TE; structural heterogeneity and graft integration. Bioprinting is highly versatile and allows for the fabrication of tissues with predetermined structure, shape, and cellularity based on custom fabricated CAD models. It is a technically demanding approach that requires multidisciplinary teams and sophisticated equipment however it can develop a productive modality of obtaining personalized solutions to a large variety of clinical demands. In situ bio printing can be envisaged as a modality to improve graft customization and host tissue integration. This will need to resolving graft maturation and mechanical stimulation within the “living bioreactor”- the mature joint environment. SF self-assembly and self-organizing allow the production of shaped structures that can be engineered to dimensionally and functionally fit implantation within the host tissue. Scaffold-less technology requires a large number of starting cells and several steps of cells/forming tissue manipulation before implantation. While this can be complicating the workflow, it allows for automation and robotization that facilitates standardization of the technological process. Enzymatic treatment or support mediated delivery are tested as modalities of resolving grafts integration within the host. Improved knowledge of developmental cascades triggered during the self-aggregation of non-differentiated cells in defined conditions could eventually lead to developing strategies for obtaining grafts of a higher level of complexity. It has been proposed that multicellular aggregates of stem cells form complex systems that can self organize in a similar manner to the one naturally occurring organogenesis. Deciphering “cytosystem dynamics” could lead to enhancing knowledge about embryonic development, tissue formation and architecture empowering novel strategies for tissue engineering [65]. Eventually, large osteochondral grafts or even entire joint complex could be engineered by streamlining self-assembly of embryonic, induced pluripotent stem cells or even adult stem cells [66].

Emerging bottom top approaches in cartilage engineering combined with the more classical top bottom techniques are speeding up the use of cartilage engineering products in everyday clinical applications.




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