Functional tissue engineering of articular cartilage for biological joint resurfacing – The 2021 Elizabeth Winston Lanier Kappa Delta Award

Abstract

Biological resurfacing of entire articular surfaces represents a challenging strategy for treatment of cartilage degeneration that occurs in osteoarthritis. Not only does this approach require anatomically sized and functional engineered cartilage, but the inflammatory environment within an arthritic joint may also inhibit chondrogenesis and induce degradation of native and engineered cartilage. Here we present the culmination of multiple avenues of interdisciplinary research leading to the development and testing of a bioartificial cartilage for tissue-engineered resurfacing of the hip joint. The work is based on a novel three-dimensional weaving (3D) technology that is infiltrated with specific bioinductive materials and/or genetically-engineered stem cells. A variety of design approaches have been tested in vitro, showing biomimetic cartilage-like properties as well as the capability for long-term tunable and inducible drug-delivery. Importantly, these cartilage constructs have the potential to provide mechanical functionality immediately upon implantation, as they will need to replace a majority, if not the entire joint surface in order to restore function. To date, these approaches have shown excellent preclinical success in a variety of animal studies, including the resurfacing of a large osteochondral defect in the canine hip, and are now well-poised for clinical translation.

Introduction

Under normal physiologic circumstances, articular cartilage functions for decades as a nearly frictionless surface in diarthrodial joints while exposed to loads of several times body weight¹. This remarkable mechanical function is attributed to the unique structure and composition of the cartilage extracellular matrix (ECM) ². In a healthy joint, the compressive, tensile, and viscoelastic properties of hyaline cartilage contribute to load bearing, energy dissipation, and joint lubrication over the lifetime of the joint ³. However, degeneration of the cartilage is associated with significant loss of cartilage function that contributes to further degeneration of the joint, which ultimately leads to osteoarthritis (OA), a debilitating disease affecting over 27 million people in the United States alone ⁴. For patients suffering from end-stage OA of the hip, the standard surgical treatment is total hip arthroplasty (THA), which has proven effective in the aging population. However, only a low percentage of young, active patients opt for THA due to the shortened projected lifespan of a hip implant for an active patient and the subsequent need for revision surgery, which is associated with significant complications, co-morbidities, overall decreased effectiveness, and less patient satisfaction ^5–7. The ability to repair or regenerate cartilage using tissue engineering strategies could have a tremendous impact on the treatment of OA for the growing population of active patients with hip OA. To this end, there has been a significant increase in research and development aimed at improving cartilage repair strategies, which include marrow stimulation, osteochondral transfer, and various forms of autologous chondrocyte implantation. However, these alternative biologic techniques are ineffective for larger lesions ^8–12. Second generation autologous chondrocyte implantation (ACI) or matrix-assisted chondrocyte implantation (MACI) procedures are not approved for use outside the knee joint and report an almost 30% failure rate within 9 years in young patients ^9,11,13. Clearly, there is a need for improved techniques, implants, and procedures to effectively treat the disease or, at a minimum, delay the progression of OA to the point at which the patients are better candidates for a hip arthroplasty procedure.

An important challenge in the development of biological resurfacing techniques for treating OA is the ability to manufacture large engineered tissue constructs with patient-specific geometries that precisely match the native joint surface, while withstanding the harsh mechanical and biochemical environment of the damaged joint. There have been several initial studies aimed at biologic cartilage joint resurfacing ^14–17. Hung and colleagues demonstrated proof of concept for joint resurfacing using young bovine chondrocytes encapsulated in agarose cultured on bovine trabecular bone, and this group later modeled nutrient and diffusion related parameters for growing large-scale constructs ^16,17. In other studies, 3D printed scaffolds were used to recreate joint anatomy in an in vivo rabbit model ^14,15, showing encouraging data demonstrating that polymer-based porous scaffolds could promote cell attachment and tissue regeneration ^14,15,18. Scaffold-free constructs have also been used in a rabbit model for articular cartilage repair, in which autologous chondrocytes are expanded in vitro to form a neo-cartilage layer ¹⁹. Condensed mesenchymal cell bodies can also be fused together to grow centimeter-sized pieces of human articular cartilage in culture ²⁰. However, the challenge of providing an implant with biomimetic cartilage properties at the time of initial cell seeding still remains and is needed to address the important clinical need for repair or regeneration of cartilage in patients who have activity-limiting cartilage loss or osteoarthritis but are too young for a total joint replacement.

Here we summarize the culmination of over 15 years of research leading to the development of a bioartificial cartilage for repair or resurfacing of large osteochondral defects, even approaching a biologically-based total joint resurfacing procedure (Figure 1). The primary technology that has served as the basis for this approach has been the development of a novel method for three-dimensional (3D) weaving of microscale biomaterial fibers, allowing the formation of tough yet flexible scaffolds for tissue engineering that allow cellular infiltration but have highly defined anisotropic, nonlinear, and viscoelastic material properties that mimic those of native cartilage. Importantly, the 3D nature of this method allows the formation of large, anatomically-shaped scaffolds that can be used for regeneration of osteochondral tissues in the shape of an entire joint surface. Here, we describe the development and refinement of this 3D weaving method for cartilage and bone tissue engineering, showing excellent functional success in a large animal preclinical study of hip OA.

Figure 1. Overview of treatment paradigm for treat advanced osteochondral disease in the young patient.

The overarching goal of this work has been to develop a tissue-engineering approach for biological resurfacing of the hip by using autologous stem cells (which may or may not be genetically altered ex vivo) and then grown on high-performance 3D woven scaffolds to create anatomically-shaped cartilage or osteochondral constructs for joint resurfacing.

Three-dimensional weaving of biomaterial scaffolds

The engineered repair of cartilaginous tissues has remained particularly challenging. From a biomechanical standpoint, cartilage can be represented as a multiphasic fiber-reinforced material, with anisotropic, inhomogeneous, nonlinear, and viscoelastic properties ^21–23. The fundamental basis for our technology for biological joint resurfacing has been through the development of a method for 3D weaving of fibers into a biomimetic scaffold that reproduces these functional mechanical properties of articular cartilage ²⁴. Previous approaches for cartilage tissue engineering have tended to utilize either porous nonwoven scaffolds ^25–30, hydrogels ^31–37, or a combination of the two³⁸ that can provide environments supportive of chondrogenesis, but generally cannot provide the complex mechanical properties believed necessary for sustained load support in vivo ²². Furthermore, the mechanical properties of most hydrogel scaffolds, particularly stiffness and strength, are several orders of magnitude lower than those of native cartilage ^28,35,39,40, thus requiring prolonged in vitro culture to attain native tissue properties before implantation. One approach to developing a scaffold with prescribed mechanical properties while maintaining an environment conducive to cell growth is by using a porous fiber-reinforced composite material that includes separate phases ^41–43.

Our primary design goal was to create a novel scaffold for the functional tissue engineering of articular cartilage that qualitatively and quantitatively mimics the behavior and mechanical properties of articular cartilage a priori, without the need for extended in vitro culture. To this end, a microscale 3D textile manufacturing technique was developed to weave bioresorbable yarns into an orthotropic, porous fabric that could be infiltrated with cells and/or a consolidating hydrogel to yield a composite scaffold (Figure 2). In contrast to standard 2D weaving, which requires lamination of separate layers to achieve the appropriate thickness, 3D weaving involves simultaneous weaving of multiple fiber layers in three orthogonal directions to form a one-piece scaffold structure with regular, interconnected pores. Additional advantages include control of multi-directional (anisotropic) mechanical properties, control of fiber spacing and volume fraction in each axis, and the ability to select each individual fiber in the construct ⁴⁴. These design variables provide for a wide range of physical and mechanical properties that can be precisely tuned to match specific tissue characteristics. By altering several of the initial design variables (e.g., material, yarn size, and yarn spacing), a composite scaffold was designed and fabricated with initial mechanical properties that were anisotropic, nonlinear, and viscoelastic, with values of mechanical test parameters that bracketed native articular cartilage a priori, even in the absence of cells and ECM.

Figure 2. Three-dimensional weaving of biomaterial scaffolds.

The basis of this technology has been the development and construction of a true 3D weaving system. (A) A custom-built 3D loom was designed and built by the authors, allowing simultaneous weaving in x, y, and z directions of 600+ resorbable fibers. (B) The primary architecture used in this work has been a macroporous orthogonal structure. Schematic representation showing x-direction (or warp) fibers in green, y-direction (or weft) fibers in yellow, and z-direction fibers in blue. (C) Surface SEMs of two 3D woven scaffolds produced with different materials and fiber volume fractions (left: PGA fibers; right: PCL fibers). (D) Cross-sections in x (top) and y (bottom) directions showing uniformity of the fabric and pores (PGA fibers). Figure adapted from ^24,47.

Initially, 3D textile structures were produced using multi-filament poly(glycolic acid) (PGA) yarn, containing rectangular pores with dimensions of approximately ~400 μm x ~300 μm x ~100 μm and a large void volume of approximately 70–75%. Composite materials were then formed by infusing agarose or fibrin hydrogels into the woven structures using a modified vacuum-assisted molding process. These composite scaffolds exhibited unique mechanical behaviors that were anisotropic, nonlinear, and viscoelastic, thereby mimicking the mechanical behavior of native cartilage. These findings show that a scaffold made from different biocompatible materials previously shown to be conducive to chondrogenesis (e.g., PGA ⁴⁵, agarose ³², or PGA and fibrin ³⁸) can be constructed with highly controlled biomechanical properties by virtue of the 3D fiber-reinforcement of the composite structure.

Tissue-engineering of cartilage using 3D woven scaffolds

To examine the ability of these 3D woven scaffolds to support cellular infiltration, cartilage matrix accumulation, and maintenance of biomimetic mechanical properties, we produced 3D woven poly (ε-caprolactone) (PCL, molecular weight 50 – 60 kDa) structures, seeded them with human adipose-derived stem cells (ASCs), suspended in a fibrin hydrogel, and cultured them for 28 days in chondrogenic culture conditions ^46,47 (Figure 3). PCL was used due to its excellent biocompatibility and slow degradation rates in vivo as compared to PGA, which degrades in 3–4 weeks. Compressive and shear biomechanical testing showed that PCL-based constructs had biomimetic mechanical properties similar to those of native human cartilage at time zero, and they maintained their mechanical properties relative to baseline throughout the culture period, while supporting the synthesis of a collagen-rich extracellular matrix. These findings are in contrast to previous studies that show culture times of 9 weeks or more are needed to achieve near-native cartilage mechanical properties ⁴⁸. Furthermore, constructs displayed an equilibrium coefficient of friction similar to that of native articular cartilage (μ_eq ~ 0.1 – 0.3) over the prescribed culture period. These findings show proof of concept for the ability of a bioresorbable and biocompatible 3D woven matrix to serve as a functional scaffold for cartilage tissue engineering. We further examined the ability of different cell sources and scaffold structures on the formation of cartilaginous tissues, as well as the ability of bioreactor-based stimulation to enhance these chondrogenic properties ⁴⁹.

Figure 3. Tissue-engineering of cartilage using 3D woven scaffolds.

(A-B) Macroscopic images of 1cm diameter implant before seeding with human MSCs and after 4 weeks of culture. (C) Compressive aggregate modulus (H_A) of implants seeded with human ASCs and cultured over 4 weeks reveal physiological compressive properties on the order of articular cartilage (dashed line shows typical mean ± s.d. (shaded region) H_A value for human cartilage ²¹. The addition of a hydrogel, fibrin, in this case, enhances the development of functional properties. (D) Type II collagen development within the pores structure of the implant is pronounced with and without the addition of fibrin with ASCs cultured on the woven PCL. (E) S-GAG biosynthesis of MSC cultured constructs was enhanced with the use of a bioreactor, and, in dual label-experiments for types I and II collagen, only type II collagen production was observed via fluorescent labeling. Figure adapted from ^47,50.

Osteochondral tissue formation on 3D woven scaffolds

There has been extensive interest in tissue-engineering cartilage, bone, or combined osteochondral constructs that can provide enhanced fixation of engineered tissues into the defect site. Thus, an important aspect of our work has been to demonstrate the ability of 3D woven scaffolds to support osteogenesis and bone formation, both in vitro and in vivo ^50–52. We hypothesized that in vitro culture duration and medium additives can individually and interactively influence the structure, composition, mechanical, and molecular properties of tissues that have been engineered with human MSCs cultured on 3D woven PCL ⁵⁰. Bone marrow MSCs were suspended in a type I collagen gel, seeded on scaffolds, and cultured for 1, 21, or 45 days in chondrogenic and/or osteogenic conditions. Structure, composition, biomechanical properties, and gene expression were analyzed. In chondrogenic medium, cartilaginous tissue formed by day 21, and hypertrophic mineralization was observed in the newly formed extracellular matrix at the interface with the underlying scaffold by day 45. Hydroxyproline, S-GAGs, calcium content, and alkaline phosphatase activity depended on culture duration and medium additives, with significant interactive effects (all p<0.0001). The 45-day constructs exhibited mechanical properties on the order of magnitude of native articular cartilage. Gene expression was characteristic of chondrogenesis and endochondral bone formation, with sequential regulation of SOX9, COL2A1, ACAN, RUNX2, BMP2, and BSP. These findings demonstrated the ability of 3D woven scaffolds to support either MSC chondrogenesis or osteogenesis dependent upon specific culture conditions.

Several previous approaches have demonstrated the ability to develop osteochondral constructs by combining multiple cells types ^53–57, multilayered scaffolds ^58–62, or multistep differentiation protocols ^54,63–69. However, it is still a major challenge to differentially direct cell fate determination into distinct lineages (i.e., cartilage and bone) from a single cell source, in a single culture system, while utilizing only one scaffold material. If proven efficacious, a single stage approach would streamline the engineering of multiphase tissues by circumventing the need for multiple cell types or multiple differentiation culture conditions. In this regard, members of the transforming growth factor β (TGF-β) family have been extensively utilized in the engineering of skeletal tissues, but it is important to note that these factors have distinct effects on chondrogenic and osteogenic differentiation of progenitor cells. We developed a method to direct human MSCs toward either an osteogenic lineage that exhibits matrix mineralization or a chondrogenesis within the same biochemical environment by culturing cells on engineered 3D woven PCL scaffolds in a chondrogenic environment while inhibiting TGF-β3 signaling through SMAD3 knockdown, in combination with overexpressing RUNX2, a master transcription factor in osteoblast differentiation ⁵¹. The highest levels of mineral deposition and alkaline phosphatase activity were observed on scaffolds with genetically engineered MSCs and exhibited an additive effect in response to SMAD3 knockdown and RUNX2 expression. Meanwhile, unmodified MSCs on PCL scaffolds exhibited accumulation of an extracellular matrix rich in S-GAGs and collagen II in the same chondrogenic environment, as expected. This ability to induce differential matrix deposition in a single culture condition opens new avenues for developing complex tissue replacements for cartilage and bone defects.

Finally, to determine if such chondral or osteochondral constructs developed in vitro can maintain their phenotype and stability in vivo, we determined the ability of MSCs to form cartilage or bone tissue, both in a nude rat subcutaneous pouch model and under simulated conditions in vitro ⁵². In the first portion of this study, various scaffold permutations, including PCL alone, PCL-bone, “point-of-care” seeded MSC-PCL-bone, and chondrogenically pre-cultured Ch-MSC-PCL-bone constructs were implanted in a dorsal, ectopic pouch in a nude rat (Figure 4). After eight weeks, only cells in the Ch-MSC-PCL constructs exhibited both chondrogenic and osteogenic gene expression profiles. Notably, while both tissue profiles were present, constructs that had been chondrogenically pre-cultured prior to implantation showed a loss of GAGs as well as the presence of mineralization along with the formation of trabecula-like structures. Notably, while both tissue profiles were present, constructs that had been chondrogenically pre-cultured prior to implantation showed a loss of GAGs as well as the presence of mineralization along with the formation of trabecula-like structures, indicating a transition to an osteogenic phenotype of MSCs following in vivo implantation.

Figure 4. Osteochondral tissue formation on 3D woven scaffolds.

Schematic of experimental setup for ectopic rat study (left) and histological results of the implant disposition at implantation (t=0 weeks) and after 8 weeks (t=8 weeks) in vivo (right). Magnified images (bottom) demonstrate high cellular activity and integration with vital bone (black arrows). Scale bars represent = 200 μm. Figure adapted from ⁵².

Hybrid composite materials made from 3D woven scaffolds

One of the major advantages of this 3D woven structure is that its high porosity and permeability, which results from a regular, interconnected pore network, allow the uniform infiltration of different materials into the structure, along with living cells. This type of approach provides many advantages such as the delivery of bioactive scaffold materials to enhance tissue growth, or the use of consolidating materials such as hydrogels or polymers to further enhance the mechanical properties of the construct^70–73. Using this approach, we have performed several studies that show that various composite scaffold such as this can not only provide fiber-reinforcement to greatly enhance the properties of chondrogenic hydrogels and materials, but the combination of complex, engineered materials can further enhance the toughness and frictional properties of the 3D woven scaffold.

In previous studies, we have shown that porous scaffolds produced from devitalized, full-thickness porcine cartilage-derived matrix (CDM) can promote the chondrogenic differentiation of seeded ASCs, without the need for exogenously added growth factors ^74,75. However, constructs showed relatively low compressive moduli (~50 kPa at day 0), and cell-seeded scaffolds contracted significantly over time. Based on these findings, we hypothesized that CDM could be combined with a 3D woven PCL reinforcement to form a functional, bioactive scaffold system capable of inducing a cartilaginous phenotype in ASCs ⁷² (Figure 5). 3D woven PCL scaffolds were infiltrated with a slurry of homogenized porcine CDM, seeded with human ASCs, and cultured for up to 42 days in standard growth conditions. While all scaffolds promoted a chondrogenic phenotype of the ASCs, CDM-only scaffolds showed low compressive and shear moduli and contracted significantly during culture. Fiber-reinforced CDM scaffolds and 3D woven PCL scaffolds maintained their mechanical properties and initial size and shape throughout the culture period, while supporting the accumulation of a cartilaginous extracellular matrix. These findings show that fiber-reinforced hybrid scaffolds can be produced with biomimetic mechanical properties, as well as the ability to promote ASC differentiation and chondrogenesis in vitro. Similarly, we have shown that infiltrating the 3D woven PCL scaffold with a self-assembling peptide (RAD16-I, also known as “Puramatrix”) can promote the re-differentiation of dedifferentiated chondrocytes and maintain the initial shape and viscoelastic behavior throughout the culture period, while constructs with RAD16-I scaffold alone contract during culture time into a stiffer and compacted structure.⁷³.

Figure 5. Hybrid composite materials made from 3D woven scaffolds.

(A) Gross morphology of cultured cartilage derived matrix (CDM), PCL, and fiber reinforced CDM (FR-CDM) constructs (6mm diam). PCL and FR-CDM constructs maintained their original size over 42 days of culture, while CDM alone showed marked contraction by day 14. (B,C) Scanning electron microscopy (SEM) surface views and (D) a cross-section of the FR-CDM scaffolds showing CDM within pores of the woven PCL scaffold (arrows). (E-F) Further demonstration of the ability of cells to infiltrate the entire scaffold as indicated by calcein AM labeling (green). (G) Immunolabeling for type II collagen demonstrates abundant type II collagen production throughout the scaffold after 6 weeks of culture. (H-K) SEM and a 3D optical profile of a 3D woven PCL infiltrated with an interpenetrating network (IPN) of alginate and polyacrylamide (Alg/PAAm). These composite surfaces demonstrate the topological smoothing effect of incorporating an IPN within the 3D woven structure. Outlined area in panel H denotes total scanned area in panel I. Outlined area in panel J denotes total scanned area in panel K. (L) Aggregate modulus of IPN and single network hydrogels and their combination with the PCL scaffold to form composites demonstrate increased compressive properties in the composite for the fiber-reinforced IPN (FR-IPN) versus the composites or scaffold alone. ∗p < 0.05 for Alg/PAAm composite vs all groups. +p < 0.05 for fibrin/polyacrylamide (Fib/PAAm) hydrogel vs all groups. ∗∗p < 0.05 for Fib/PAAm composite vs Alginate or Fibrin composites or scaffold alone. Figure adapted from ^70,72.

In addition to regulating cell behavior, an infiltrating hydrogel or polymer in the 3D woven scaffold can be used to “consolidate” the fibers and significantly alter or enhance their mechanical behavior. To test this process, 3D woven scaffolds were infiltrated with various interpenetrating network (IPN) hydrogels, which consist of two different polymers that are mixed with one another at the molecular scale (Figure 5). These polymers are especially attractive as they exhibit synergistically increased fracture toughness and tribological properties as compared to the individual components of the network ⁷⁶. We found that infusion of the 3D woven PCL scaffold with IPNs of alginate/polyacrylamide (Alg/PAAm) or fibrin/PAAm significantly increased the Young’s modulus over the PCL scaffold alone ⁷⁰. Specifically, the aggregate modulus of the Alg/PAAm IPN (~0.4 MPa) was significantly higher than that of the single network hydrogel of alginate or PAAm separately. However, the reinforcement of the hydrogel with the 3D woven scaffold resulted in a further improvement in the modulus of the composite construct, with values reaching approximately 1.2 MPa for 3D woven PCL infiltrated with Alg/PAAm IPN and 0.9 MPa for the 3D woven PCL infiltrated with fibrin/PAAm IPNs, which were significantly higher than the PCL scaffold alone (Figure 5). This finding suggests that the interaction of the IPN hydrogels with the 3D woven PCL scaffold was responsible for the improvement in mechanical strength, a departure from conventional interpretation of fiber reinforced gel networks, in which the fiber components generally serve as the strength determinant ⁷⁰.

In follow-up studies to this work, we developed an agarose and poly(ethylene) glycol IPN hydrogel that showed high viability of MSCs within the IPN hydrogel, with improved mechanical properties compared to constructs comprised of individual components ⁷¹. We further strengthened these properties by integrating the hydrogel with a 3D woven structure. The resulting fiber-reinforced hydrogels displayed functional macroscopic mechanical properties mimicking those of native articular cartilage, while providing a local microenvironment that supports cellular viability and function. These findings suggest that a fiber-reinforced IPN hydrogel can support stem cell chondrogenesis while allowing for significantly enhanced, complex mechanical properties at multiple scales as compared to individual hydrogel or fiber components.

Scaffold-mediated viral delivery for spatial and temporal control of cell behavior

While many tissue-engineering approaches are based on the delivery of stem or progenitor cells on scaffolds, the differentiation of these cells can require extensive in vitro manipulation and costly growth factor delivery. Furthermore, upon implantation in vivo, the ability to further specify or modify cell fate is lost. Thus, the ability to induce and maintain stem cell differentiation in vivo over extended periods of time in the absence of exogenous growth factors could significantly enhance tissue regeneration. To address this issue, we developed a method to functionally immobilize lentiviral particles precisely to individual fibers of the 3D woven scaffold using poly-L-lysine (PLL) (Figure 6) ^51,77,78. This approach allows site-specific transduction of cells that are seeded on the scaffold, while preventing the virus from affecting other cells.

Figure 6. Scaffold-mediated viral delivery for spatial and temporal control of cell behavior.

(Top Row) Tunable expression in MSCs with a dox-inducible lentiviral vector. (A) Schematic representation of immobilized lentivirus on PCL fiber. (B) Schematic diagram of lentiviral vectors with TGF-β (top) and IL-1Ra (bottom) as the representative gene of interest. (Middle Row) Scaffold-mediated LV transduction of human MSCs within 3D woven PCL scaffolds (C) SEM image of a 3D woven PCL scaffold 5 mm disk (scale bar = 1mm) and fluorescence image of constitutive eGFP-expressing MSCs on the 3D woven PCL. (D) Scaffold-mediated lentiviral delivery was used to transduce MSCs with the gene for TGF-β, resulting in the production of tissue with a cartilaginous phenotype as demonstrated by Safranin-O red and fast green staining for GAGs and collagen, respectively, without the addition of exogenous TGF-β. (Bottom Row) Tunable IL-1Ra expression in engineered cartilage constructs (E) IL-1Ra secretion from engineered cartilage constructs into media every 72 hours over 36 days of chondrogenesis (mean ± SEM, n=3). +Dox indicates dox induction at 1 μg/mL. -Dox indicates baseline IL-1Ra expression. Arrows indicates when dox (1 μg/mL) was switched on (↑) and off (↓) every 9 days. (F) Safranin-O red and fast green staining showing loss of S-GAGs in Control (non-transduced) cells and protection from IL-1 by IL-1Ra-producing cells. The white space shows the location of the PCL fiber bundles. Figure adapted from ^77,78.

Scaffold-mediated transduction of MSCs with lentiviral vectors driving expression of TGF-β3 led to potent chondrogenic differentiation and accumulation of a cartilaginous ECM rich in type II collagen and S-GAGs, both of which are major components of articular cartilage ⁷⁷. By nearly all measures of gene regulation, protein content, and biomechanical properties, the level of chondrogenic differentiation achieved at day 28 by immobilizing lentivirus was indistinguishable from that using exogenous TGF-β3 treatment ⁷⁷.

In addition to delivering genes for the production and directed regulation of anabolic growth factors ^51,77, we have also used this approach to create tunable and inducible systems for delivery of cytokine inhibitors ⁷⁸. For example, there is ample evidence indicating that proinflammatory cytokines, and particularly interleukin 1 (IL-1), play an important role in the pathogenesis of OA ^79–81 as well as the inhibition of mesenchymal stem cell (MSC)-based repair of cartilage ^82–87. While we have demonstrated that the overall mechanical functionality of cartilage engineered tissue in the presence of pathophysiological levels of IL-1⁸⁸ can be preserved through the use of a 3D woven scaffold, exposure to this cytokine significantly inhibits the chondrogenic development and maturation of MSC-synthesized ECM ⁸⁹. As a means of inhibiting the effects of such an inflammatory environment on engineered cartilage, we used this lentivirus-immobilization technique to create a tunable and inducible gene delivery system for IL-1 receptor antagonist (IL-1Ra)⁷⁸, the inhibitor of IL-1 (Figure 6). A doxycycline-inducible vector was used to transduce MSCs within the 3D woven PCL scaffolds to enable tunable IL-1Ra production. In the presence of IL-1, IL-1Ra-expressing engineered cartilage produced cartilage-specific extracellular matrix, while resisting IL-1-induced upregulation of matrix metalloproteinases and maintaining mechanical properties similar to native articular cartilage ⁷⁸. The ability of functional engineered cartilage to self-tune the delivery of anti-inflammatory cytokines to the joint may enhance the long-term success of therapies for cartilage injuries or osteoarthritis.

Anatomically shaped tissue-engineered cartilage for biological joint resurfacing

Biological resurfacing of entire articular surfaces represents a novel but challenging strategy for treatment of cartilage degeneration that occurs in osteoarthritis. Not only does this approach require anatomically sized and functional engineered cartilage, but the inflammatory environment within an arthritic joint may also inhibit chondrogenesis and induce degradation of native and engineered cartilage. To address these issues, we used adult stem cells to engineer anatomically shaped, functional cartilage constructs capable of tunable and inducible expression of anti-inflammatory molecules, specifically IL-1Ra ⁹⁰. Large (22 mm diameter) hemispherical scaffolds were fabricated from 3D woven PCL fibers into two different configurations and were seeded with human ASCs (Figure 7). Doxycycline(dox)-inducible lentiviral vectors containing eGFP or IL-1Ra transgenes were immobilized to the PCL to transduce ASCs upon seeding, and constructs were cultured in chondrogenic conditions for 28 days.

Figure 7. Anatomically shaped tissue-engineered cartilage for biological joint resurfacing of the hip (top) and femoral condyle of the knee (bottom).

(left) Lesions and joint anatomy are mapped from MRI (or CT) and used to prepare an anatomically shaped mold of the hip or knee. (middle) 3D scaffolds were woven from PCL fibers. The implant is formed in a mold, matching anatomical contours of the articular surfaces, and seeded with human ASCs (right). Tissue-engineered hip resurfacing implant culture after 5 weeks of in vitro culture. Figure adapted from ⁹⁰.

Constructs showed biomimetic cartilage properties and uniform tissue growth while maintaining their anatomic shape throughout culture. IL-1Ra-expressing constructs produced nearly 1 μg/mL of IL-1Ra upon controlled induction with dox. Treatment with IL-1 significantly increased MMP activity in the conditioned media of eGFP-expressing constructs but not in IL-1Ra-expressing constructs ⁹⁰. Our findings show that advanced textile manufacturing combined with scaffold-mediated gene delivery can be used to tissue-engineer large anatomically shaped cartilage constructs, in the shape of a hip or knee condyle, that possess controlled delivery of anti-cytokine therapy. Moreover, these cartilage constructs have the potential to provide mechanical functionality immediately upon implantation, a critically important characteristic, as they will need to replace a majority, if not the entire joint surface, in order to restore function.

Biological resurfacing in a large animal, preclinical model of hip osteoarthritis

Given the success and proof-of-concept in all aspects of this work in vitro and in small animal orthotopic sites, we next sought to test the ability of 3D woven, tissue-engineered osteochondral construct to resurface a massive defect on the femoral head of the hip and to restore the biomechanical function of the tissue and the whole joint ⁹¹. An important step in this process was the development of a bicomponent scaffold consisting of the 3D woven textile described in the previous section, bonded to a 3D printed osteoconductive PCL base for direct fixation into the femoral head. Autologous bone marrow derived mesenchymal stem cells (MSCs) were harvested, expanded in culture, seeded on the scaffold, and differentiated using a chondrogenic cocktail containing TGF-β3 for 4 weeks (Figure 8). To minimize potential regulatory issues, gene therapy approaches were not used in this preclinical study. A 10 mm in diameter x 2 mm deep osteochondral lesion was created on the dorsal, load bearing aspect of the femoral head (where canine hip OA typically develops) in skeletally-mature purpose-bred hounds and repaired with the cell-seeded implant matching the anatomical joint surface curvature. For control animals, this defect was left empty. The implants showed success in all measured outcomes. The dogs that received the engineered implant returned to normal activity levels by 6 months as measured by their activity during the day, averaged over a 2-week period, whereas the control cohort did not return to normal activity levels by the end of the study. Kinetic gait analysis revealed normal GRFs (peak vertical force [PVF]; vertical impulse [VI]) on the operated limb by 6 months postoperatively in the implant group. Conversely, the control animals displayed clinically observable lameness throughout the study after OA was initiated in the joint and significantly lower GRFs than the implant group. The pain index score, which measures behavior, activity levels, happiness- and anxiety-like behaviors, demonstrated a return to normal behavior by 6 months in the implant group, in direct contrast to the OA control animals.

Figure 8.

Biological resurfacing in a large animal, preclinical model of hip osteoarthritis. (Top Row) Schematic of implant design. (A-B) Textile component is fused to the (C) additive manufactured component to form the (D) biocomponent implant of variable heights. (Middle row) Disposition of the implant at the time of implantation. (E) Gross image of implant (10mm diameter) that was seeded with autologously harvested canine MSCs and cultured in vitro for 18 days in chondrogenic conditions. Note smooth, cartilaginous tissue synthesized in upper surface layer. (F) Confocal image demonstrating viable ECM synthesis (fluorescent labeling for collagen) is confined to upper layer of bilayered implant (scale bar = 0.5 mm). (G) Intraoperative photograph of massive defect created in the femoral head, and (H) image immediately after placing the implant. (Bottom row) (I) Representative examples at necropsy of operated femoral heads at 6 months (arrow indicates location of implant). Safranin O and Fast Green (left) and Trichrome stain (right). Note that unstained section of each image is structural polymer. (J) Modified O’Driscoll scores and (K) elastic compressive equilibrium (Young’s) modulus demonstrate a significant tissue repair (Groups not sharing the same letter are statistically different *p<0.05). Figure adapted from ⁹¹.

All implants remained firmly implanted in the osteochondral defect and demonstrated consistent integration with the surrounding cartilage with no evidence of inflammation. In contrast to control defects, the tissue-engineered implants exhibited a smooth surface that matched the gross morphology of the joint, and no macroscopic depressions were noted in the histological sections. Mechanically, the compressive modulus of the repaired cartilage was 10-fold higher than the tissue from control defects and similar to that reported for normal canine hip cartilage. By all outcome metrics – which included activity monitoring, gait/standing analysis, muscle circumference, and behavioral indices – animals receiving the implant returned to normal preoperative values by 6 months postop. Our data, in total, indicate that this engineered implant restored the contour of the femoral head to its native condition and was functioning as intended from a structural, biological, and anatomical viewpoint to restore joint function and relieve pain in a model of hip OA.

Conclusions and Future Directions

The culmination of this research supports the utility of this approach for large-scale cartilage repair in the clinical setting. A primary advantage of this approach is due to the ability to precisely tune the multidirectional mechanical properties during scaffold manufacture, prior to cell seeding, while maintaining a large void fraction to allow cellular infiltration and tissue ingrowth. Further, 3D-orthogonally woven fabrics have the ability to conform to anatomical structures without compromising the mechanical or biological performance of the structure. The orthogonal interconnected pore structure of both the 3D woven scaffold and the bicomponent scaffold utilized in the canine study enable direct and open communication with the surrounding native tissues, enabling solid integration with host tissues. These properties have also proven highly beneficial in the development of novel cell-based implants for that combine synthetic biology with tissue-engineering for autonomous drug delivery capabilities⁹². Future directions are focused on the translation of this technology into clinical trials for the treatment of osteochondral lesions in the hips young patients (i.e., less than 60 years old) for whom there are currently no good therapeutic options, with the goals of further extending this technology to other joints (e.g., knee, ankle, shoulder, etc.). From a technological standpoint, the combination of 3D weaving with 3D printing provides a platform technology that allows for production of anatomically-shaped bicomponent implants with tunable mechanical and biological properties, and we hope to continue to develop new applications of this technology for the treatment of various musculoskeletal conditions.