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3D BIOMIMETIC, BI-PHASIC KEY FEATURED SCAFFOLD FOR

OSTEOCHONDRAL REPAIR

Cross-reference to Prior Applications

[0001] This application claims priority to US provisional patent application 61/799,243 filed on March 15, 2013, the entire contents of which are hereby incorporated by reference.

U.S. Government Support

[0002] This invention was not made with support from the United States Government.

Background of the Invention

Area of the Art

[0003] The present invention relates to a novel method for the creation of three dimensional (3D) biologically inspired tissue engineered scaffolds with both excellent interfacial mechanical properties, and biocompatibility. In some embodiments, a combination of nanomaterials, nano/microfabrication methods and 3D printing can be employed to create structures that promote tissue reconstruction and/or production. In other embodiments, electrospinning techniques can be used to create structures made of polymers and nanotubes.

Description of the Background Art

[0004] Osteochondral defects as a result of trauma, congenital, and/or pathological disorders present a crucial clinical problem [1 , 2, 3], Osteochondral defects penetrate the entire thickness of articular cartilage, beyond the calcified zone, and into the subchondral bone. The osteochondral tissue is a nanostructured tissue notoriously difficult to regenerate due to its extremely poor inherent regenerative capacity, complex stratified architecture and disparate biomechanical properties [1, 3]. Although various biomaterials and tissue engineering approaches to treat osteochondral defects have been investigated, it is still very challenging to replicate the robust integration of the cartilage and subchondral bone and the complex stratified cartilage/bone structure. None of the current available treatment options provides a perfect solution for osteochondral regeneration.

[0005] As modern medicine advances, novel methodologies are being explored and developed in order to solve and improve current cartilage and osteochondral problems [4, 5, 6, 7], In particular, two approaches that can be used to create integrated scaffolds are electrospinning and 3D printing. Regarding electrospinning, the use of several novel techniques has made it possible to modify the properties of generated scaffolds. Co- electrospinning and wet electrospinning, for example, have proven to be very useful techniques for the generation of complex scaffolds. In addition, the use of polymers, polysaccharides and inorganic extracellular matrix (ECM) components have led to scaffolds with enhanced mechanical characteristics. However, the improvements made are still not sufficient to successfully create extremely complex scaffolds that can replicate complex tissues such as cartilage or the bone-cartilage interface.

[0006] 3D printing is emerging as a complex tissue manufacturing technique, and offers great precision to control the internal architecture of a scaffold and print complicated structures close in architecture to native tissue [8]. More importantly, based on computer- aided design (CAD) models, 3D printers can fabricate a predesigned patient-specific tissue construct in a layer-by-layer fashion [9, 10, 11, 12, 13]. Furthermore, non-invasive MRI images of patients' osteochondral defects, can be obtained and used to inform CAD design, which would allow the scaffold to perfectly fit into the defect site and be ideal for the patient specific shape required critical sized osteochondral implant. Recently, Cui et al. successfully Inkjet printed a poly(ethylene glycol) dimethacrylate solution containing chondrocytes into a defect formed in an osteochondral plug [10]. They observed greater proteoglycan deposition in the interface of implant and native tissue. Current attempts, while producing viable 3D tissue scaffolds, still lack higher sophistication both in the ability to control and define osteochondral scaffold micro architecture.

[0007] Recently, 3D printing and rapid prototyping processes have been used to create scaffolds that are 3D with user defined micro-structures and micro-scaled architectures [10, 14]. This ensures that the scaffold not only is fully unoccluded with uniformly interconnected pores, but also that a great many more complex, predesigned architectures patterns and structures can be implemented. Hard tissue is one of the most readily researched and treated defect and injury sites for Tissue Engineering (TE) scaffold-based solutions. One of the critical 3D scaffold design criteria for hard tissues is that they must have suitable mechanical properties. In addition, interconnected pores, specifically pore structures at the micro-scale, interconnected by smaller pores on a nano- scale are also indicative of the ECM of hard tissues, and are very important for hard tissue scaffold design [15, 16, 17]. This sort of complicated, hierarchical structure is one that is difficult to replicate, if at all, and then is more difficult to control in even very advanced electrospinning setups and other common scaffold fabrication techniques. With the application of 3D printing, there is an allowance not only for the creation of delicate and intricate structures from the advanced working of strong and robust materials, but the potential to create highly ordered structures that could conceivably match any desired architecture [18]. This later advantage is one that also makes 3D printing attractive for other types of targeted tissue 3D scaffolds.

[0008] A method that is very popular for 3D printing of joint tissue is fused deposition modeling (FDM). Fused Deposition Modeling (FDM) is one of the simplest forms of 3D fabrication. In FDM, a computer-aided design (CAD) drawing is used in conjunction with a 3D printer to create polymeric 3D structures. A FDM machine consists of a slightly heated printing bed, a printing head capable of 3D axial movement and a computer/controller. The printing head draws a solid polymeric filament from a spool and forces it through a heated extruder head, which heats the material and deposits it as a thin, molten layer on the printing surface. The machine then prints multiple thin layers on top of the previously deposited layer. In the end, one is left with a 3D construct of pre-determined design [19, 20]. FDM is somewhat rudimentary as compared to other 3D fabrication methods, but it is important because it establishes an overarching methodology in all 3D fabrication techniques, where a fully 3D-designed structure is disassembled into very thin, successive slices and then physically recreated layer-by-layer. FDM itself has strong potential as a 3D fabrication method for 3D TE scaffolds because of its ability to employ a number of different polymers but is not often utilized because it lacks a high enough resolution to create complex and biomimetic nano/microstructures [21].

[0009] In an example of cutting edge 3D printing for multi-tissue systems, Shim et al. used a deposition system similar to FDM called solid freeform fabrication. A 3D scaffold was printed from a deposited, structurally sound polymer, while a cell laden hydrogel was infused into the void space of the printed structure. The printed hard scaffold served as a structural support while the printed soft hydrogel served to encapsulate cells and ensure their even distribution throughout the printed construct [22].

[0010] A challenge and unmet need in the art is the creation of 3D printed osteochondral scaffolds with both excellent interfacial mechanical properties and biocompatibility for facilitating human bone marrow mesenchymal stem cell (MSC) differentiation. Such a scaffold would need to be designed to have special mechanical considerations. Previously, work exploring the osteochondral regeneration has yielded scaffolds that are weak at the interface between the cartilage and bone regions. Often, s caffolds are fabricated in two or three layers separately and then joined together with a glue or suture [23, 24].

Summary of the Invention

[0011] The present invention relates to a novel method for the creation of 3D biologically inspired tissue engineered scaffolds with both, excellent interfacial mechanical properties, and biocompatibility. In some embodiments, a combination of nanomaterials, - - nano/microfabrication methods and 3D printing can be employed to create structures that promote tissue reconstruction and/or production.

[0012] In exemplary embodiments, nanomaterials and nano/microfabrication methods are used to create novel biologically inspired tissue engineered cartilage scaffolds for facilitating MSC chondrogenesis. The methods disclosed herein can be readily adapted by someone of ordinary skill in the art to create a variety of tissue scaffolds that can be used for the promotion and generation of tissue repair and/or production. In this sense, tissue can be interpreted as biological material that is made up of epithelial cells, muscle cells, connective tissue cells, nerve cells and/or blood cells. In exemplary embodiments, electrospinning and/or 3D printing techniques can be used to design a series of novel 3D biomimetic nanostructured scaffolds based on carbon nanotubes and biocompatible poly(L-lactic acid) (PLLA) polymers. Polylactic acid, variously known as poly(lactic) acid, polylactide, PLA or PLLA, is a biocompatible and biodegradable thermoplastic polymer. It consists of an aliphatic polyester of L-lactide units. These various names for polylactic acid are used interchangeably herein. Specifically, a series of electrospun fibrous PLLA scaffolds with controlled fiber dimension can be fabricated. These fabricated scaffolds can promote attachment of MSCs as can be shown by in vitro MSC studies in which the stem cells prefer to attach in the scaffolds with smaller fiber diameter. Additionally, in some embodiments, these scaffolds can be incorporated with biomimetic carbon nanotubes and ploy-L-lysine to induce more chondrogenic differentiations of MSCs.

[0013] In other exemplary embodiments, 3D printed polymer constructs can be generated using the methods disclosed herein. In some embodiments, these 3D polymer constructs can be designed to mimic the certain tissues and/or organs, including the osteochondral region of the articulate joint, and to have enhanced mechanical characteristics when compared to traditional bi-phasic designs. In some embodiments, these fabricated 3D printed polymer constructs can be subjected to surface modification, both with a chemically functionalized acetylated collagen coating and through absorption via poly-L-lysine coated carbon nanotubes so as to promote the growth and differentiation of MSCs.

[0014] One of ordinary skill in the art can recognize that the use of the techniques and methods described herein can be applied towards the generation of 3D printed polymer constructs that mimic a variety of tissues and/or biological environments and are not necessarily restricted to cartilage tissue engineering applications as the embodiments disclose. In addition, one of ordinary skill in the art can appreciate that these constructs can also be modified to include surface modifications (or other modifications not exclusive to the surface) that can more appropriately mimic the native tissue or environment with which they are intended to interact. In addition, one of ordinary skill in the art can readily appreciate that these constructs can be further modified to more specifically and/or efficiently promote the differentiation, growth, and/or production of cells and tissues specific to a particular biological environment and/or organ.

[0015] Further objectives and advantages, as well as the structure and function of preferred embodiments will become apparent from a consideration of the description, and non-limiting examples that follow.

Description of the Figures

[0016] FIGURE 1 shows scanning electron microscopy (SEM) images of electrospun fibers. A-E show electrospun fibers spun at a working distance of 12, 14, 16, 18 and 20 cm (respectively). 1 F represents a lower magnification of electrospun fibers at 18cm.

[0017] FIGURE 2 is a bar graph showing MSC attachment to PLLA scaffold as a function of fiber diameter.

[0018] FIGURE 3 shows SEM and TEM images of H ₂ untreated or treated multi-walled carbon nanotubes (MWCNTs). Figure 3A is a SEM of MWCNTs, 3B is a SEM of H ₂ treated MWCNTs, 3C is a TEM of MWCNTs, 3D is a TEM of H ₂ treated MWCNTs.

[0019] FIGURE 4 shows SEM images of varying electrospun fibrous PLLA scaffolds. Figure 4A is of pure PLLA scaffold prepared by normal dry electrospinning, 4B is of pure PLLA scaffold prepared by wet electrospinning, 4C is of 1% H2 treated MWCNTs and 4D is of 0.5% MWCNTs in PLLA.

[0020] FIGURE 5 is a bar graph of enhanced Young's Modulus of MWCNT in electrospun PLLA scaffolds as compared to controls.

[0021] FIGURE 6 is a bar graph of MSC proliferation on various electrospun/MWCNT scaffolds.

[0022] FIGURE 7 is a bar graph of glycosaminoglycan (GAG) synthesis of MSCs in all MWCNT embedded in PLLA scaffolds.

[0023] FIGURE 8 is a bar graph of total collagen synthesis of MSCs in H ₂ treated MWCNT embedded PLLA scaffolds.

[0024] FIGURE 9 shows CAD images of large and small pore models. 9A shows large pore models, 9B shows small pore models. 9C is a picture 3D printed scaffolds with different internal geometry and pore density in cell growth media. Small (top) and large (bottom) pore models, and (from left to right) homogeneous, bi-phasic and bi-phasic key scaffolds.

[0025] FIGURE 10 is a bar graph of Young's Modulus data for 3D printed scaffolds. [0026] FIGURE 11 is a bar graph of shear fracture energy of 3D printed scaffolds, performed under wedge shear fracture shear testing.

[0027] FIGURE 12 is a collection of SEM images of uncoated and acetylated collagen type I coated 3D printed PLA scaffolds. 12A-C show uncoated PLA scaffolds, and 12 D-F show coated PLA scaffolds.

[0028] FIGURE 13 is a bar graph of MSC proliferation in a variety of 3D printed PLLA scaffolds with different internal structure and surface modification.

[0029] FIGURE 14 is a bar graph of GAG synthesis in various 3D printed osteochondral scaffolds.

[0030] FIGURE 15 is a bar graph of collagen type II synthesis on 3D printed scaffolds.

[0031] FIGURE 16 is a bar graph of total protein synthesis.

[0032] FIGURE 17 shows 3D models of a full knee cartilage layer model and a 3D printed cartilage scaffold. 17A represents the 3D models; and 17B is an image of the printed scaffold.

[0033] FIGURE 18 is a bar graph of MSC proliferation in a variety of 3D printed PLA scaffolds with different internal structure and surface modification.

[0034] FIGURE 19 shows 3D images of a 3D scaffold as viewed from different angles. 19A is a 3D image of the structure as transparent and as viewed from the top, 19B and 19D are images of the structure (transparent or not) as viewed from an elevated side angle; 19C represents an image of the structure as transparent as viewed from the side. .

Detailed Description of the Invention

[0035] The following description is provided to enable any person skilled in the art to make and use the invention and sets forth the best modes contemplated by the inventor of carrying out his invention. Various modifications, however, will remain readily apparent to those skilled in the art, since the general principles of the present invention have been defined herein specifically to provide sample embodiments of the invention. In describing embodiments, specific terminology is employed for the sake of clarity. However, the invention is not to be limited to the specific terminology so selected. While specific exemplary embodiments are discussed, it should be understood that it is for illustration purposes only. A person skilled in the art will recognize that other components and configurations can be used without parting from the spirit and scope of the invention. All references cited herein are incorporated by reference as if each had been individually incorporated to the extent permitted by applicable law and regulation. - -

[0036] While other fabrication methods have been widely and thoroughly investigated for scaffold fabrication, they still present a number of limitations such as having weak or poor mechanical properties, having non-uniform pore distribution, random pore interconnectivity and void space, and limited control over the size, and distribution of fibers within the micro and nano architecture of the scaffold [25, 26], Electrospinning has been used as a method to generate electrospun scaffolds, and although it comes with some caveats, it presents a system amiable to manipulation. Recently, 3D printing and rapid prototyping processes have also been used to create scaffolds that are 3D with user defined micro-structures and micro-scaled architectures [27, 28]. This ensures not only that the scaffold is fully unoccluded with uniformly interconnected pores, but a great deal of more complex, predesigned architectures patterns and structures can be implemented.

[0037] Regarding electrospinning, several parameters have been extensively studied in modulating properties of electrospun scaffolds to include: the choice of polymer, polymer concentration, working distance, and voltage [29]. The effects on cell behavior of altering the physical and chemical properties of electrospun scaffolds have been extensively studied on osteoblasts and chondrocytes. The current landscape of bone and cartilage regeneration research via electrospun scaffolds has focused on novel electrospinning techniques, and the employment of biomimetic composite materials for enhanced cell function, as well as directed stem cell proliferation and differentiation through chemical modification of fabricated scaffolds. With regards to hard tissue scaffolds, modification of scaffold physical properties without compromising mechanical integrity is of great concern.

[0038] A number of innovative methods for enhancing electrospun scaffold characteristics have been investigated. One of the most widely researched areas being co-electrospinning [30-38]. This has been shown to be a highly versatile option when attempting to fabricate a scaffold with desirable characteristics. Another advanced method for electrospun scaffold fabrication is wet electrospinning where fibers are collected in a solvent bath. This method has been shown to create highly porous materials with complex, interconnected pore structure rendering them ideal for bone regeneration. Aside from variations in fabrication techniques, novel nanomaterials such as DNA-based self- assembled nanotubes have also been used in electrospun nanocomposites.

[0039] Traditional materials, such as natural polymers, polysaccharides and inorganic extracellular matrix (ECM) components have been used extensively via incorporation within polymeric scaffolds in an effort to enhance the mechanical characteristics of scaffolds, as well as improve cell behavior. Although progress has been made, the improvements made are still not ideal prompting researchers to investigate novel - - materials, such as nano diamond crystals (ND-OCTs) and rosette nanotubes (RNTs) with attractive and unique qualities. Increasingly, researchers have begun to turn to unconventional, unique materials to improve the functionality of electrospun scaffolds beyond what has been capable with conventionally applied materials for electrospun biomimetic nanocomposite scaffolds. In addition to the use of new materials, the methods for which scaffolds are fabricated can also be modified. Research is moving in the direction of developing new, more complex electrospinning methods which have the potential to yield more complicated and characteristic architectures (i.e., vascularized electrospun scaffold).

[0040] 3D printing provides an alternative to electrospinning for scaffold formation. Hard tissue is one of the most readily researched and treated defect and injury sites for tissue engineering scaffold based solutions. One of the critical 3D scaffold design criteria for hard tissues is that they must have suitable mechanical properties. In addition, interconnected pores, specifically pore structures at the micro-scale, interconnected by smaller pores on a nano-scale are also indicative of the ECM of hard tissues, and are very important for hard tissue scaffold design. This sort of complicated, hierarchical structure is one that is difficult to recapitulate, if at all, and then more difficult to control in even very advanced electrospinning setups and other common scaffold fabrication techniques. With the advent of 3D printing, there is a possibility not only for the creation of delicate and intricate structures from the advanced working of strong and robust materials, but a potential to create highly ordered structures that could conceivably match any desired architecture [2]. This later advantage is one that also makes 3D printing attractive for other types of targeted tissue 3D scaffolds.

[0041] Currently, 3D printing as applied to TE uses a layered manufacturing method of printing thin depositions of material in a given pattern on top of previously printed and cured material [2, 28], This could allow for large, macro-scale objects that have complex, user-defined internal features, mimicking the architecture a given organ. This could also allow for materials to be printed which encapsulate living cells into the artificial organ construct, creating a complex network of cells with an advantageous architecture conducive to organ function and cell/tissue growth [39].

[0042] Moreover, one of the most important challenges facing 3D TE construct design is vascularization. Scaffolds seeded with cells that begin to mature and form tissue have problems with the transportation of nutrients and essential signaling chemicals and growth factors, as well as removal of waste products within the internal structure of the scaffold [31 , 40, 41]. In the body, vascular networks accomplish these tasks, but new and under- formed vasculature present a daunting limitation to scaffold-based tissue repairs. - -

However, if a scaffold can be fabricated with designed transport channels and structures that mimic vascularized tissue, then it could be possible to ameliorate this limitation [42]. 3D printing presents a potential ability to accomplish this because, as stated previously, it is possible to create structures with predesigned complex, micro-scale internal architectures.

[0043] Currently in the field, several unique fabrication methods for controlled, 3D TE scaffolds have been recently investigated. 3D fiber deposition is similar to FDM, where a heated nozzle is used to deposit a melted polymer, but the outlet used is on the order of several hundred micrometers in diameter. The process yields micro-fibrous structures, and the spacing and deposition angle of fibers can be modulated. Fedorovich et al used 3D Fiber Deposition to create alginate hydrogel matrices containing chondrocytes and osteogenic progenitors, as well as separate printed layers for osteoblasts and osteoblast growth for osteochondral defects. Good cellular growth results were reported and a high degree of effect was demonstrated on the scaffold architecture by modulation of the above mentioned process parameters [43], Sun et al. also used 3D fiber deposition to create and compare porous PCL scaffolds containing osteoblasts, which were fabricated at 45 degree and 90 degree deposition angles. The 3D printed scaffolds were compared to traditional salt-leeched scaffolds. The cell distribution on the 3D scaffolds was more homogeneous than the salt-leached scaffolds, demonstrating that 3D scaffolds are more effective for tissue engineering. The results also showed that it is possible to design and optimize the properties of amorphous polymer scaffolds by 3D fiber deposition [44].

[0044] Lu et al. also utilized projection printing which works similarly to photolithography, where a photo-mask is used to cure layers of photosensitive material in designed patterns when exposed to light. In projection printing, a UV light source is used in conjunction with a micro-mirrorarray, a digital masking device, imaging optics and a photocurable resin to photopolymerize the resin into complex, biomimetic shapes. Lu et al. was able to use this process to print precise closed channels and cavities that mimicked native vasculature [45].

[0045] Relevant to this invention, Shim et al. used a deposition system similar to FDM called solid freeform fabrication. A 3D scaffold was printed from a deposited, structurally sound polymer, while a cell laden hydrogel was infused into the void space of the printed structure. The printed hard scaffold served as a structural support while the printed soft hydrogel served to encapsulate cells and ensure their even distribution throughout the printed construct [46]. [0046] Other methods of 3D printing used include, but are not limited to: selective laser sintering [47-55], laminated object manufacturing [56, 57], and inkjet 3D printing [14, 57, 58].

[0047] The abovementioned examples of electrospinning methods and 3D printing methods are not intended to be a comprehensive overview of all methods in the art. Also, the examples listed below are not to limit the scope of the invention. One of ordinary skill in the art could take the invention disclosed herein and modify it to better replicate a particular tissue or biological environment through the use of different polymer fibers, and/or incorporation of specific factors, tissues, cells, biological factors such as DNA, RNA, peptides, or other chemical agents. A variety of entities can be used to deliver additional compounds/factors such as micro/nano spheres, tubes or fibers through diffusion or other means.

[0048] Example 1 : Electrospun nano/microscaffold for cartilage tissue engineering

[0049] The purpose of these experiments was to investigate if the mechanical and cytocompatibility properties of electrospun polymer scaffolds for cartilage repair could be enhanced, with the addition of nanomaterials. It was also a goal to evaluate if the nanotubes modified with a cell-favorable molecule can effectively control specific differentiation of stem cells.

[0050] Advances in tissue engineering require more sophisticated materials both to characterize and grow tissues. For this purpose, carbon nanotubes/fibers are emerging candidates. Although the use of carbon nanotubes in tissue engineering is at its infancy, they have been considered exciting alternatives as templates for tissue growth, drug delivery agents and in bio-sensory applications. Carbon nanotubes mimic the dimensions of the constituent components of tissues, where cells are accustomed to interact with nano-fibrous proteins. This property makes them excellent candidates for invoking positive cellular responses when employed as implants. In addition, the superior mechanical properties of carbon nanotubes are efficient for their use as a secondary phase for high load bearing applications. Their electrical properties make them a potential choice in neural applications where signal transfer between growing axons necessitates electrical conductance. The unique chemical properties they possess permit them to be functionalized with different chemical groups, which further promote cell growth.

[0051] For instance, while there have been several studies that show the ability of CNTs and their conductive properties to incite cardiac tissue development of stem cells, Serag et al, in a rather dramatic study, studied the effects of CNTs on plant cells that were differentiating into tracheary elements [59]. It was found that these cells readily used "cup- sacked" CNTs to create cell structures via oxidative cross-linking of monolignols to the CNT surface. This not only demonstrates CNTs having a desirable effect on cell growth, but also highlights potential CNT fate in a living dynamic system post-application. Furthermore, there have also been experiments, where CNTs are used in a nanocomposite material, with promising results. Ogihara et al studied the use of CNT alumina ceramic composites in vivo for bone tissue engineering. It was found that there was no increased inflammation at the implant site of CNT containing samples when compared to alumina controls, suggesting that constituent CNTs embedded in a matrix material do not necessarily exhibit some of the harmful effects of CNTs. Although there have been cytotoxicity concerns raised about CNTs, thus far the exact mechanisms of CNTs' effects on cells is still not fully known. But all of the results presented show that nanotubes are cytotoxic only under certain conditions, e.g. certain tube lengths, hydrophobicity of nanomaterial and dispersion of nanotubes. It has also been reported that the formation of IVIWCNT aggregates can significantly contribute to their inflammatory qualities [60].

[0052] Materials and Methods

[0053] Hydrogen treated and non-treated multi-walled carbon nanotubes (MWCNT)

[0054] MWCNTs were obtained from Shanghai Xinxing Chenrong Technology Development CO., LTD. The MWCNTs were synthesized by the floating-catalyst technique in chemical vapor deposition process. Dimethylbenzene (C ₈ H ₀ ) and Thiophene (C4H4S) served as carbon sources and the Iron atom from Ferrocene (Fe(C ₅ H ₅ ) ₂ ) was used as catalyst for the growth of MWCNT. The synthesis processes were carried out in a cylindrical chamber with temperature of 1100°C in hydrogen environment. In some studies, MWCNTs were hydrogen treated. Briefly, the procedure involved placing the MWCNTs in a mixture of nitrogen and hydrogen environment at a temperature of 800°C for two hours and then removing the hydrogen supply and allowing the samples to cool naturally. The hydrogen treatment removes amorphous carbon and nanohorns encapsulating metal catalyst nanoparticles and MWCNT, making tubes more uniform. The morphologies of these tubes, both treated and untreated, were evaluated using Scanning Electron Microscopy (SEM) and Transmission Electron Microscopy (TEM).

[0055] Scaffold fabrication

[0056] For all experiments, poly L-lactic acid (PLLA), purchased from Sigma Aldrich was used as the base polymer to be electrospun. PLLA comprises of polymerized L-lactide units with ester linkages between each lactide. Fibers were fabricated using an in house setup, consisting of a VWI syringe pump, Harvard Apparatus variable voltage supply and an aluminum collector plate. For the scaffold preparation for adhesion study, PLLA was dissolved at 18% weight by volume in a 9 to 1 solution of Dichloromethane (DCM) and - -

Dimethyl formamide (DMF). DMF evaporates at a higher temperature than DCM, and was thus postulated to induce porosity. All adhesion study scaffolds were electrospun at 12, 14, 16, 18 and 20 cm working distance from the collector plate, at voltages varying from 14 to 18 kilovolts (kV).

[0057] The proliferation and differentiation study used PLLA dissolved in pure DCM at 18% w eight by volume, which was then wet-electrospun into a coagulation bath of methanol. PLLA scaffold were also electrospun with solutions containing 0.5% w/v untreated MWCNTs, 0.5% w/v H ₂ treated MWCNTs and 1% w/v H ₂ treated MWCNTs. For the electrospun PLLA, MWCNTs were blended into the solvent-polymer solution and then sonicated prior to electrospinning. These coated tubes, in solution, were then added to the scaffolds, and allowed to soak in a 37°C incubator for 24 hours. This caused a simple absorption coating to form.

[0058] All samples were mixed using ultrasonication, with the nanotubes being sonicated in the solvent first in order to assure uniform distribution.

[0059] Scaffold Characterization

[0060] Microscopy was done on samples coated with gold nanoparticles, which were then viewed using a Zeiss SigmaVP Scanning Electron Microscope (SEM). Transmission Electron Microscope (TEM) images were taken with a JEOL JEM-1200EX TEM. When electrospun into PLLA fibrous scaffolds, the dispersion of each species of nanotubes was also evaluated using TEM microscopy. Fibers were coated in an approximately 4-8 nm of gold nanoparticles to make them image able and then placed on copper grids to facilitate imaging.

[0061] Mechanical testing

[0062] All mechanical tests were done using and ATS axial tester, a 50 Newton load cell and compression placard. Samples from each of the experimental groups were also mechanically tested. Circular samples of 8 millimeters in diameter and about 1 to 2 mm in height were taken and tested in compression, at a strain rate of 0.2 mm per second. The force-deformation data was then used to calculate and compare the Young's Modulus of each sample.

[0063] MSC cell culture

[0064] Primary human bone marrow MSCs were derived from healthy consenting donors from the Texas A&M Health Science Center, Institute for Regenerative Medicine and thoroughly characterized. They were used to evaluate the cytocompatibility properties of the nanocomposite coatings. MSCs (passage #3-6) were cultured in a complete media comprised of Alpha Minimum Essential medium (a-MEM, Gibco, Grand Island, NY) - - supplemented with 16.5% fetal bovine serum (Atlanta Biologicals, Lawrenceville, GA),1% (v/v) L-Glutamine (Invitrogen, Carlsbad, CA), and 1% penicillin : streptomycin solution (Invitrogen, Carlsbad, CA) and cultured under standard cell culture conditions (37°C, a humidified, 5% C02/95% air environment). For the differentiation study, chondrogenic media was prepared which consisted of the above media recipe, but with the addition of 100nM dexamethasone, 40 pg/ml proline, 100 pg/ml sodium pyruvate, 50 g/ml L- Ascorbic acid 2-phosphate and ITS+ at a concentration of 1% of the total volume of prepared media.

[0065] MSC adhesion study

[0066] Sample groups consisting of dry-spun pure PLLA scaffold, fabricated at working distances of 12, 14, 16, 18, and 20 cm were then seeded with MSCs at 10,000 cells per scaffold and allowed to incubate in a biological incubator for 4 hours. Cells were then lifted and counted using a hemocytometer and light microscope. The experiment was repeated three times.

[0067] MSC proliferation study in vitro

[0068] Sample groups consisting of a pure PLLA control, PLLA scaffold containing 0.5% w/v untreated MWCNTs, PLLA scaffold containing 0.5% w/v H ₂ treated MWCNTs and PLLA scaffold containing 1.0% w/v H ₂ treated MWCNTs were then seeded with MSCs at 10,000 cells per scaffold. All sample groups were fabricated using wet-electrospinning. Samples were then incubated and collected at one day, three days and five days. Cells at each day were lifted with Trypsin EDTA, reacted with Thermoscientific photometric cell counting reagent (MTS assay) and analyzed using a Thermo Scientific Multiskan GO microplate reader at a setting of 490 nm wavelength light.

[0069] MSC differentiation study

[0070] Finally, a differentiation study was conducted, using a pure PLLA wet- electrospun scaffold and four more experimental groups containing 0.5% w/v MWCNTs, 0.5% w/v H ₂ treated MWCNTs, 0.5% w/v MWCNTs coated with poly-L-lysine and 0.5% w/v H ₂ treated MWCNTs coated in poly-L-lysine. The poly-L-lysine was used as a means to increase the hydrophilicity of the scaffolds and test the efficacy of MWCNTs dispersed in a polymer scaffold as a chemical delivery device. Samples were cultured in chondrogenic media with MSCs seeded at 250,000 cells per scaffold, and incubated for two weeks, with samples being collected at one and two weeks. Collected samples were freeze dried in a lyophilizer and treated in a Papain digestion solution for the chondrogenic differentiation evaluations. [0071] Glycosaminoglycan (GAG): Glycosaminoglycan, a key component of cartilage, was measured using a standard BlyscanTM GAG assay kit. Samples were centrifuged in a microcentrifuge of 10 minutes at 10,000 RPM and then reacted with a dye reagent for 20 minutes. Samples were then centrifuged again, reacted with a dissociation reagent and analyzed in the microplate reader at 560 nm.

[0072] Total collagen synthesis: Total collagen content of the samples was also evaluated using a Sircoll collagen standard assay kit. Samples were prepared in the same way as the GAG assay, except that before the addition and removal of dye, the collagen was reacted with a coagulation reagent and refrigerated overnight. Samples were also evaluated with the microplate reader at 560 nm.

[0073] Total protein content: Total protein content in the cell lysates was measured using a commercial BCATM Protein Assay Reagent kit (Pierce Biotechnology) and following manufacturer's instructions. For this purpose, 150 lit of aliquot supernatants of the protein- containing cell lysates were mixed with 150 1AL of working agent solutions (including1 :24:25 of cupric sulfate : bicinchoninic acid : a reagent with sodium bicarbonate, sodium carbonate and sodium tartrate) and then were incubated at 37°C for 2h. Light absorbance was measured at 562 nm on the spectrophotometer. According to a standard curve of known concentrations of albumin versus absorbance run in parallel with experimental samples, the total protein synthesized by MSCs cultured on the substrates of interest to this study was calculated.

[0074] MSC proliferation study in vitro:

[0075] All cellular studies used human bone marrow derived MSCs cultured in cell culture media as described previously. Once all the scaffolds were fabricated and modified, a proliferation study was conducted, with constructs being seeded at 200,000 cells per scaffold and were cultured in a MSC growth media for 1 , 3 and 5 days. After the prescribed time periods, adherent cells were quantified via a MTS assay as described above.

[0076] Statistics:

[0077] All cellular experiments were run in triplicate and repeated three times for each substrate. Data are presented as the mean value ± standard error of the mean (SEM) and were analyzed with student's t-test for pair-wise comparison. Statistical significance was considered at p < 0.05.

[0078] Results and Discussion

[0079] Electrospun nanostructured scaffold for cartilage regeneration - -

[0080] The adhesion study sought to compare the effectiveness of fiber size on cellular adhesion, as well as to optimize our setup. As shown in Table 1 , smaller fiber diameters were yielded by increasing the working distance, with a slight increase at 20 cm. These results can also be observed via SEM imaging (Figure 1). In Figure 1 , A-E represent electrospun fibers at 12, 14, 16, 18 and 20 cm (respectively). Figure 1F represents a lower magnification of electrospun fibers at 18cm.

[0081] Table 1 : Electrospun fibers' diameters, as compared to working distance between the needle and the collector plate

[0082] More importantly, it is shown that the scaffolds with the smallest fiber diameter promoted the greatest cellular adhesion of MSCs (Figure 2), displaying that the smallest and thus more biomimetic fiber dimensions promote the best stem cell adhesion. Here, we see significantly enhanced MSC attachment on the electrospun PLLA scaffold with the smallest fiber diameter (1.33 micrometers). It should be noted that our larger fibers also provided good cell adhesion. This may be due to the fact that the larger fibers had induced nanoscale (60 nm) surface pores, which may have helped to create more surface area for MSC adhesion. This suggested both of fiber dimensions and surface topography can contribute to create a biomimetic stem cell-favorable environment.

[0083] MWCNT PLLA scaffolds, MSC proliferation and chondrogenic differentiation in vitro

[0084] SEM and TEM images of the untreated MWCNTs compared to the H ₂ treated MWCNTs were taken (Figure 3). Figure 3A is a SEM of MWCNTs, 3B is a SEM of H ₂ treated MWCNTs, Figure 3C is a TEM of MWCNTs, 3D is a TEM of H ₂ treated MWCNTs. These images show H ₂ heating changed morphology of the nanotubes from bundles of nanotube aggregates into more homogenous distribution, which make them suitable for co-electrospinning into PLLA scaffold. In addition, the H ₂ treatment can facilitate removing impurities and metallic catalyst material in nanotubes. SEM images were also taken of the electrospun PLLA scaffolds fabricated via dry and wet electrospinning and with MWCNTs (Figure 4). Figure 4A is of pure PLLA scaffold prepared by normal dry electrospinning, 4B is of pure PLLA scaffold prepared by wet electrospinning, 4C is of 1% H2 treated MWCNTs and 4D is of 0.5% MWCNTs in PLLA. The wet electrospun scaffold shows more 3D porous structure than dry electrospun scaffolds. No nanotubes could be observed, implying that they were fully imbedded within the fibers. It could also be seen that the fiber diameters varied on the MWCNT embedded scaffolds. The fiber diameter distribution was from about 1 micrometer to 10 micrometers.

[0085] More importantly, after the addition of H ₂ treated or untreated MWCNTs, the PLLA scaffolds' Young's modulus increased dramatically when compared to a pure PLLA control (Figure 5, *p<0.05, ** p<0.05), and all MWCNT reinforced scaffolds were within the range of native articulate cartilage (-0.75 to 1 MPa). Figure 5 shows the significantly enhanced compressive Young's Modulus of MWCNTs in electrospun PLLA scaffolds as compared to PLLA controls. This shows that the incorporation of just a small amount of MWCNTs can increase the mechanical properties of a tissue engineering scaffold to within biomimetic regimes.

[0086] The proliferation study showed an increase in cellular proliferation on all scaffolds, with the greatest cell numbers on the scaffolds with incorporated MWCNTs of both species at three days (Figure 6, n=9, *p,0.05, **p,0.05). Figure 6 shows MSC proliferation on various electrospun PLLA MWCNT scaffolds. At five days all of the scaffolds showed even greater cellular growth. Both of nanotube species may change the nanosurface roughness and surface area of the scaffolds, thus contributing to the significantly improved MSC proliferation after 5 days when compared to controls. The differentiation study also showed increased chondrogenic differentiation activity (Figures 7 and 8). Figure 7 shows GAG synthesis of MSCs in all MWCNT 9.5%) embedded PLLA scaffolds as compared to controls. There was a dramatic increase in GAG content at one and two weeks on the scaffolds containing poly-L-lysine coated MWCNTs, and among those samples the H ₂ treated tubes performed the best (Figure 7) (#, *, **, *** p<0.05). The positively charged poly-L-lysine can create an electrostatic interaction with negative charged GAG for improved GAG nucleation in the scaffold. This implies that the surface coating of the nanotubes (to decrease hydrophobicity) had the greatest impact on GAG synthesis in vitro, far greater than the nano surface topography contribution of MWCNTs. Furthermore, Figure 8 reveals that both H ₂ treated MWCNT and poly-L-lysine coated MWCNT PLLA scaffolds can significantly improve total collagen synthesis after one and two weeks (*p<0.05, ** p<0.05, # p<0.05, ## p<0.05, ### p<0.05, $ p<0.05). Specifically, Figure 8 shows improved total collagen synthesis of MSCs in H2 treated MWCNT embedded PLLA scaffolds. The fact that the H ₂ treated tubes yielded the better results shows that the purification of nanotubes to remove various impurities and modify - - nanotube morphology is advantageous for biological applications. It should be noted that the scaffold containing untreated nanotubes didn't improve collagen synthesis when compared to the control. This is, however, not wholly unexpected, since it is known that the metal catalyst material and impurities have cytotoxic of the addition of CNTs to the construct.

[0087] Poly-L-lysine has been long-established as a beneficial chemical compound for promoting cellular growth, and recently has been used in regenerative studies. Grothman et al. investigated the secondary structure of peptide chains and their effects on proliferating osteoblasts. It was found that the peptides in Poly-L-lysine adopt an intermolecular beta sheet structure. This reveals an increased area of spread, which consequently supports osteoblast proliferation. In addition, Santana et al used poly-L- lysine coated slides to culture human chromaffin progenitor cells [61]. The coated sliders were able to induce the cells to differentiate into two distinct neuron-like cell types.

[0088] In our study, we implemented a small concentration of homogenously distributed MWCNTs which have been H ₂ purified and poly-L-lysine coated for the chondrogenic differentiation of MSCs for the first time. Our results show the significant beneficial effects of these nanostructured scaffolds in directing stem cell differentiation in vitro. We believe that our scaffolds are advantages for cellular growth because the nanotubes are modified to have cell-favorable hydrophilic poly-L-lysine coated surfaces, are homogenously dispersed and imbedded in solid microfibrous structures, which creates a stable and advantageous environment for cellular activity.

[0089] Example 2: 3D printed scaffolds for osteochondral regeneration

[0090] Materials and Methods

[0091] 3D osteochondral scaffold design and fabrication

[0092] All 3D osteochondral scaffolds were designed using Rhinoceros 3D modeling package. Scaffolds were then printed in groups of six using a PrinterBot 3D printing system, modified with a 347 pm diameter nozzle, and a spool (or filament fed into the printer) of 1.75 mm diameter biocompatible Polylactic acid (PLA) polymer. PLA comprises an aliphatic polyester of L-lactide units. 3D models were converted into a gcode instruction file using Slic3r, and then used to instruct the printer via the Pronterface software package. The PLA was extruded into a filament using a screw extrusion method. Raw polymer, usually in the form of small beads or pellets, is fed into a hopper attached to a heated, tubular chamber. The chamber has a motor driven screw throughout which turns and moves the pellets down the chamber, and melting them along the way. The chamber terminates at a nozzle, where the melted plastic is forced through to form a filament. The filament, while still soft, is usually collected on a spool. Filaments were purchased from MakerBot.

[0093] There were a total of six experimental groups designed (Figure 9). The first group was a homogenous cross-hatched structure (Figures 9A and 9B, top panels). The second was a bi-phasic structure consisting of a cross hatched pattern (110) and an intersecting rings structure (111) (Figures 9A and 9B middle panels). The hatch pattern, which has already been widely used for 3D printed joint repair, was chosen for both its proven performance as a biomimetic micro pattern and as a biomimetic analog to the alignment of ECM and chondrocytes in articulate cartilage. The ring structure for the bone layer was chosen as a means for designing randomly oriented, interconnected pores, which mimics the structure of subchondral bone. Finally, a biphasic key model (Figures 9A and 9B, bottom panels) with an internal structural feature (112) in the form of a cylinder or multiple connected cylinders traversing the length of the scaffold, was designed and printed. All scaffolds described were designed as cylindrical plugs, 14 mm in diameter and 8 mm high. All of the above models have both small and large pore features (500 micrometers and 1000 micrometers, respectively) based on the printing limitations of the setup at the time. Each layer in the "bone region" of the scaffolds shifts about 400 micrometers, so that each layer is staggered, forming a fully interconnected porous network. Also, for the small-featured, bi-phasic key scaffold, the inner radius of the cylindrical tube encasing the internal features of the construct is 0.5 cm and the outer radius is 1 cm and 1 mm. For later shear testing, we also designed an additional intermediate pore size (750 micrometers) for the three models. These pore sizes and features, when printed and analyzed under SEM, were virtually the same in size as those of the 3D models originally designed. In addition, a full knee model with anatomical shape was designed and printed.

[0094] Moreover, we applied a collagen type I coating on the printed scaffolds (i.e., bi- phasic key osteochondral scaffold with small pore feature) to further improve their cytocompatibility properties. A protocol for chemically functionalized attachment known as acetylation was utilized. Type I Collagen which had been pre-acetylated was purchased. Briefly, scaffolds were immersed in an ethylenediamine/n-propanol (1 :9 ratio) solution at 60°C for 5 min. They were then extensively washed with deionized water and dried at 35°C. The aminolysed scaffolds were then immersed in a 1% glutaraldehyde solution at room temperature for 3 h to transform the NH ₂ groups into CHO groups. After washing extensively, the scaffolds were immersed in 0.1% acetylated collagen at 4 °C for 24 h. The process itself yields a series of layered chemical attachments, finally resulting in a collagen coating. From the PLA up, we have an ester linkage between the PLA and the ethylenediamine, a "Schiffs base" linkage between the ethylenediamine and the glutaraldehyde and further Schiffs base linkage between the glutaraldehyde and the collagen.

[0095] As opposed to a chemical process, hydrogen-treated MWCNT used in the previous example were also attached to scaffolds using absorption. A solution of 0.1% poly-L-lysine dissolved in de-ionized water was added to dry MWCNTs and ultrasonicated for 90 minutes. This solution was then added to dry scaffolds, 1 ml per scaffold, and incubated overnight. The samples were then removed from solution after 24 hours, washed in de-ionized water and dried at 60°C.

[0096] Mechanical testing, modeling and scaffold imaging

[0097] All mechanical testing including compressive and shear testing was conducted using a uniaxial testing system (ATS systems). For compression testing, a flat 2 cm in diameter platen was attached to a 500 N load cell. The platen was then advanced into the scaffolds, oriented uniaxially with the bone layer on the bottom and the cartilage layer interacting with the platen, at 0.02 cm/min. Data were taken using LabView, and then analyzed in Microsoft Excel. Load and displacement was used to plot the stress / strain curves and then Young's modulus was taken from the liner elastic region. For shear testing, the same setup and conditions were used, with the exception of the platen being replaced with a 5° wedge (from centerline, 10° total) and the scaffold rotated 90°. The interface between the bone and cartilage layers was aligned parallel to the wedge, and the wedge was advanced into the interface line for bi-phasic and key scaffolds. For homogeneous models, the wedge was advanced into the scaffold at half of the scaffolds' height, which is consistent to the dimensions and orientations of the other two models. Force was plotted against displacement and the area under the curve was taken to provide the shear fracture energy in N/mm ² .

[0098] Based on the obtained experimental data, a computational model was established to estimate and correlate the properties of various structures with different porosities. In addition, a Zeiss SigmaVP Scanning Electron Microscope (SEM) was used to image the surfaces of acetylated collagen constructs and controls (uncoated scaffolds). Scaffolds were coated with an approximately 4-8 nm of gold nanoparticles and then isolated on carbon tape dots to facilitate imaging.

[0099] In vitro MSC evaluation

[00100] Cell culture: Primary human bone marrow MSCs were derived from healthy consenting donors from the Texas A&M Health Science Center and thoroughly characterized. They were used to evaluate the cytocompatibility properties of the 3D printed scaffolds. Details are provided in Example 1 , above. They were subsequently lifted from samples for analysis using Trypsin-EDTA.

[00101] MSC proliferation: Once all the scaffolds (six osteochondral models and one biphasic key model with collagen) were fabricated and sterilized, a five day proliferation study was conducted in 24 well plates, with cells seeded at 100,000 cells per scaffold and 2 ml of media per well. Details are provided in Example 1 above.

[00102] MSC chondrogenic differentiation:

[00103] A two week differentiation study was also conducted on scaffolds with optimal pore density decided by MSC proliferation (i.e., small pore features). New scaffolds were fabricated, of the same physical specifications with small pores (control, bi-phasic and biphasic key) and an extra set of key models were coated with acetylated collagen. Chondrogenic media was prepared which consisted of the above media recipe, but with the addition of dexamethasone, proline, sodium pyruvate, L-Ascorbic acid 2-phosphate, TGF-βΙ and ITS+. MSCs were seeded at 150,000 cells per scaffold and cultured in the chondrogenic media. Samples were then taken at 1 and 2 weeks. The following standard chondrogenic biochemistry assays were used to evaluate MSC chondrogenic differentiation in our 3D printed scaffolds.

[00104] Glycosaminoglycan (GAG) content: GAG was measured using a standard GAG assay kit (Accurate Chemical & Scientific Corp., Westbury, NY) according to manufacturer's instructions.

[00105] Type II collagen synthesis: Human type II collagen was evaluated via a type II collagen ELISA assay (Fisher Scientific, Pittsburgh, PA). Briefly, control and sample aliquots were added to a precoated 96- well plate and incubated. Unbound sample was washed and a horse radish peroxidase-labeled collagen II antibody was added, incubated, and washed. After washing, tetramethylbenzidine was added producing a blue color. The reaction was stopped by the addition of an acidic stop solution and read at 450 nm.

[00106] Total protein synthesis: Total protein was evaluated using a Micro BCA assay (Thermo Scientific, Rockford, IL). An uncultured collagen coated scaffold control was also digested and tested for total protein content. This measurement was then subtracted from the weeks 1 and 2 total protein analysis.

[00107] Statistics

[00108] All experimental data was compiled as mean ± standard error mean for each property measured. Numerical data were analyzed via one-way ANOVA and student's t- test to determine differences amongst the groups. Statistical significance was considered at p<0.05.

[00109] Results

[00110] Structure and mechanical characterization of 3D printed scaffolds

[001 1] Figure 9 shows our novel cylindrical osteochondral construct design and printed scaffolds, with different internal structure. These homogenous (Figure 9A and 9B top panels) and biphasic scaffolds (Figure 9A and 9B middle and bottom panels) were designed to establish both a control group and a more traditionally designed osteochondral scaffold for comparison of key featured design (Figure 9A and 9B bottom panels). Figure 9C is a picture of the printed structures. The homogenous model is a uniformly patterned structure, mimicking only one type of tissue. The bi-phasic scaffold is more similar to traditional ostochondral scaffolds, containing both a cartilage and bone layer and no other materials or features. Each layer in the "bone layer" of the scaffolds shifts about 400 micrometers, so that each layer is staggered, forming a fully interconnected porous network. This key feature was designed to traverse the entire length of the scaffold, and penetrates both the cartilage and bone layer. It was intended to increase overall mechanical strength and to prevent failure of the device at the bi-phasic interface between the bone and cartilage layers. Physical characteristic data of all of printed scaffolds was computed from 3D models of all the scaffold groups (Table 2). It can be seen that the total surface area of the construct increases from a homogenous design to a bi-phasic design, and again when a key feature is added. Furthermore, the total surface area of the construct increases again when the feature size is decreased. However, the surface area to volume ratio of the construct follows the opposite trend as described above. With a decrease in feature size, more features can be added to the construct, thus increasing the overall volume, and is not a reflection of the surface to volume ratio of a given feature. Mechanical compression tests were also conducted on the six different scaffold construct designs (Figure 10). In Figure 10, Data are ± standard error mean, n=5; *p<0.05 when compared to all homogenous and biphasic scaffolds; **p<0.05 when compared to all other scaffolds with small features; and #p<0.05 when compared to all other scaffolds. All of the scaffolds showed excellent mechanical properties similar to or exceeding cartilage (0.75 to 1 MPa) and subchondral bone (30 to 50 Mpa) in human osteochondral tissue. Under compressive loading, the biphasic key models both in small and large feature have the highest modulus when compared to the homogeneous controls and the bi-phasic models. The bi-phasic scaffolds with large features performed better than the similar constructs with small features.

[00112] Table 2: Physical Data for Different 3D Constructs and Pore Sizes Total

Smallest Bulk

Pore Density Surface SA/V

Feature Volume

(pores/mm ^A 3) Area Ratio (mm) (mm ^A 3)

(mm ^A 2)

Large

Homogeneous 1 0.5 1850.6 616.3 3.002

0.5 bone

Bi-phasic 1 - 4 2094.4 716.2 2.924

0.001 cartilage

0.5 bone

Bi-phasic Key 1 - 4 2150.7 749.8 2.868

O.OOlcartilage

Intermediate

Homogeneous 0.71 0.53 2368.6 576.3 4.109

0.71 - 0.53 bone

Bi-phasic 2700.8 863.9 3.126

1.7 0.003 cartilage

0.71 - 0.53 bone

Bi-phasic Key 2724.6 904.3 3.012

1.7 0.003cartilage

Small

Homogeneous 0.5 5.3 2817.7 571.1 4.933

5.3 bone

Bi-phasic 0.5 - 2 2854.0 863.6 3.304

0.005cartilage

5.3 bone

Bi-phasic Key 0.5 - 2 2921.7 947.4 3.083

0.005cartilage

[00113] Shear fracture energy testing was conducted on our bi-phasic key scaffold, bi- phasic scaffold and homogeneous controls, with three varying pore sizes, for a total of nine scaffolds (Figure 11); Data are ± standard error mean, n=5; ^Λ ρ<0.01 when compared to controls with intermediate pores. In all cases, the scaffolds showed a trend in the force per unit area that it took to cleave the scaffolds apart, increasing from the homogeneous control to the bi-phasic model to our novel key model. Moreover, the surface morphology of our collagen coated scaffolds was imaged by SEM as shown in Figure 12; SEM images of uncoated (A-C) and acetylated collagen type I coated (D-F) 3D printed PLA scaffolds. These scaffolds exhibited a collagen texturing when compared to uncoated samples.

[00114] Computational modeling

[00115] The computational process has great potential for optimized 3D printing design, since is easier and faster to perform than experimental measurement. Generally, Poisson's ratio varies in a small range, so for our purposes we assume it as a constant. The mechanical properties of the newly designed porous structure were calculated through a relationship between Young's modulus and porosity, which has been widely discussed. Rossi [63] modified Hashin's equation so that Young's modulus {E) is a function of low concentration of spherical pores, i.e. ^{o )} , where ⁰ stands for the Young's modulus of the parent solid, P refers to the total porosity volume fraction, and B is a geometric parameter. Based on this, Rice [64] proposed an exponential function which can be applied for a wide range of pore character. Later, this empirical formula was successfully applied to predict the mechanical properties of porous hydroxyapatite bioceramic [65],

[00116] Improved MSC proliferation and chondrogenic differentiation in vitro

[00117] The five day proliferation study (Figure 13) showed that the biphasic scaffolds with small pore features can significantly promote MSC proliferation after 5 days; Data are ± standard error mean, n=9; *p<0.05 when compared to all other scaffolds and **p<0.05 when compared to all scaffolds with large features and homogenous controls with small features at day 5. It should be noted that all of the scaffolds experienced a decrease in cellular activity from day one to day three. Furthermore, the scaffolds with acetylated collagen outperformed all other groups, which show that the chemical modification can greatly increase MSC proliferation.

[00118] For 2 week chondrogenic differentiation, each sample was analyzed for GAG, total protein and collagen type II synthesis. Results of the GAG assay showed the most GAG deposition present on the key and collagen coated key scaffolds after one and two weeks (Figure 14); Data are ± standard error mean, n=9; &p<0.05 when compared to all other scaffolds and $p<0.05 when compared to controls after two weeks; and ^Λ ρ<0.05 when compared to controls and biphasic scaffolds after 1 week. More interestingly, all samples showed an increase, with the far greatest increase on the key scaffold, but not on the collagen coated key scaffold.

[00119] In contrast to our GAG result, all biphasic and biphasic key scaffolds with and without collagen coating showed greatly enhanced type II collagen deposition when compared to controls (Figure 15); Data are ±SEM, n=9; *p<0.05 when compared to all other scaffolds at week 1 and ^Λ p<0.05 when compared to all other scaffolds at week 2. All samples showed increased type II collagen synthesis when compared to week 1. The key model is not intended to explicitly direct differentiation by modifying the mechanical cues of the microenvironment, but rather to strengthen the bulk construct (at the interface) in a physiological environment. Figure 16 s hows i ncreased total protein content on biphasic, key scaffold with/without collagen coating after 1 week when compared to controls, with the most total protein present on the key model; Data are ± standard error mean, n=9; ^Λ ρ<0.05 when compared to controls, & p<0.05 when compared to all other scaffolds and &&p<0.05 when compared to bi-phasic and controls after two weeks. At week 2 all samples continued to increase when compared to controls, but with the largest increase on the collagen type I coated scaffolds.

[00120] Knee concept design and fabrication

[00121] In addition to samples printed for cellular study and mechanical testing, a large construct, mimicking the structure of a human knee with internal bi-phasic and key features was designed (Figure 17). Figure 17A represents 3D models while 17B is an image of the printed scaffold. This model also had superficial pores on the surface, to allow fluid perfusion in a theoretical in vivo scenario. When exposed to fluid, there was an ease of perfusion through the full construct, showing that the internal architecture was interconnected.

[00122] Conclusion

[00123] We have designed and fabricated, using CAD and 3D printing, a series of novel biocompatible scaffolds for osteochondral tissue repair. These scaffolds sought not only to recreate the cartilage and bone layers of the osteochondral region, but ultimately to incorporate special mechanical reinforcement elements, dubbed "key" features, which were intended to increase the mechanical strength and integration of the two distinct tissue zones. Mechanical testing showed that key scaffolds performed better in both compression and in shear when compared to homogeneous controls consistently across a variety of different pore sizes. These results were then further supported by computational analysis of scaffold mechanical properties. This implies that our constructs would perform better under natural mechanical loading at the osteochondral interface in situ. In addition, a MSC proliferation study and chondrogenic differentiation study were conducted. In both cases, our key scaffolds or key scaffolds with collagen coatings outperformed controls. Novel key scaffolds displayed enhanced MSC growth, and expressed more chondrogenic synthesis of GAG, type II collagen and total protein content than controls. This suggests that our key scaffolds provide a robustly integrated bone-cartilage construct that could withstand mechanical stresses post-implantation and effectively regenerate cartilage at the osteochondral interface.

[00124] Example 3: Additional 3D printed Osteochondral Devices

[00125] Other embodiments were made with features similar to those in Figure 9, including: 1) a homogenous cross-hatched structure, with features of 1 to 0.5 mm in size, 2) a bi-phasic structure consisting of a cross hatched pattern and an intersecting rings structure, and 3) biphasic structures but with reinforced key feature in the interface. In addition to above samples printed for cellular study and imaging, a large construct, mimicking the structure and anatomical shape of a human knee with internal bi-phasic and key features was also designed (similar to that shown in Figure 17). A Stratasys Fortus 250 m 3D printing system was used to fabricate the full large model out of Acrylonitrile butadiene styrene (ABS), a common material used in rapid prototyping 3D printing, for demonstration purpose. Furthermore, the 3D printed cartilage layer of the model was synthesized of biocompatible PLA polymer. This model also had superficial pores on the surface, to allow fluid perfusion in a theoretical in vivo scenario. In addition, a plain sample, a collagen coated sample and a multi-walled carbon nanotube (MWCNT) coated sample were produced using small featured bi-phasic key featured scaffolds. For the MWCNT-coated 3D printed PL.A structures, MWCNTs were sonicated in poly-L-lysine, causing a simple coating on the nanotubes. These coated tubes, in solution, were then added to the scaffolds, and allowed to soak in a 37°C incubator for 24 hours. This caused a simple absorption coating to form.

[00126] Physical characteristic data was computed from the 3D models of all of the scaffold groups (Table 3). It can be seen that the total surface area of the construct increases from a homogenous design to a bi-phasic design, and again when a key feature is added. Furthermore, the total surface area of the construct increases again when the feature size is decreased. However, the surface area to volume ratio of the construct follows the opposite trend as described above. This is due to the fact that, with a decrease in feature size, more features can be added to the construct, thus increasing the overall volume, and is not a reflection of the surface to volume ratio of a given feature.

[00127] Table 3: 3D printed scaffolds' physical characteristics based on computed 3D model data. - -

[00128] Mechanical compression tests were also conducted on the six different scaffold

[00129] construct designs, as seen previously in Figure 10. All of the scaffolds showed excellent mechanical properties similar to or exceeding cartilage (0.75 to I.O MPa) and subchondral bone (30 to 50 Mpa) in human osteochondral tissue. Under compressive loading, the biphasic key models both in s mall and large feature have the highest modulus when compared to the homogeneous controls and the bi-phasic models. The biophasic scaffolds with large features performed better than the similar constructs with small features.

[00130] Moreover, the proliferation study result (Figure 18) (*, ** p,0.05) showed on day one, with slightly greater cellular activity on bi-phasic scaffolds when compared to homogenous control groups. More importantly, our result shows that all of the biphasic scaffolds with small features can significantly promote MSC proliferation after 5 days. Based on table 3, these biphasic scaffolds with smaller feature attain increased surface area and greater feature density, thus providing a more advantageous environment for cellular growth. Furthermore, the scaffolds with acetylated collagen and poly-L-lysine coated H2 treated MWCNTs outperformed all other groups, which shows that nanostructured surface morphology and chemical modification can greatly increase MSC proliferation. SEM images of the surface topography of these surface modified scaffolds were captured as in Figure 12. It was also observed that all scaffolds showed a decreased MSC proliferation on day three, which may be due to the fact that these constructs had very large internal features, and cells have been shown to cease proliferative activity when migrating through a construct.

[00131] Example 4: Program Designed Scaffold

[00132] Description

[00133] The perfected design in Figure 19 is made up of small hexagonally shaped pores in the bone region (201), and lateral "rods" which alternate 90 degrees with each layer in the cartilage region (202), in order to form square shaped pores and highly aligned channels. The bone region has pores of 350 micrometers in diameter, and features sizes of 200 micrometers wide and 350 micrometers high. The cartilage region has pores of 150 micrometers wide, and features that are 200 micrometers wide. The gross bone region is 7 mm thick, and gross cartilage region 3 mm thick with a 200 micrometer solid superficial layer, with an overall diameter of 10 mm for the entire device. Each layer in the bone region shifts about 400 micrometers, so that each layer is staggered, forming a fully interconnected porous network. Inside the device, there are a tubular structures, dubbed "key" features, with a 600 micrometer outer radius and a 300 micrometer inner radius, which traverse the device from the bottom to the top (203), but does not penetrate the top (204) or bottom (205) layers. There are a total of 9 of these key features. Figure 19 contains 3D images of the structure as viewed from different angles: 19A represents a 3D image of the structure as transparent as viewed from the top; 19B and 19D represent images of the structure (transparent or not) as viewed from an elevated side angle; 19C represents an image of the structure as transparent as viewed from the side.

[00134] The device is f abricated via 3D printing technology, using fused deposition modeling to print polylactic acid (PLA), and/or paste deposition, biofilament plotting and stereolithography [51 , 52, 53] to print hydrogels of polyethylene glycol (PEG), polyethylene glycol diacrylate (PEG-DA) and polyethylene glycol methacrylate. Nanocrystalline hydroxyapatite (nHA) rich PLA and / or PEG are used for the bone region and "key" features, and a PEG varietal is used for the cartilage region.

[00135] The following claims are thus to be understood to include what is specifically illustrated and described above, what is conceptually equivalent, what can be obviously substituted and also what essentially incorporates the essential idea of the invention. Those skilled in the art will appreciate that various adaptations and modifications of the just-described preferred embodiment can be configured without departing from the scope of the invention. The illustrated embodiment has been set forth only for the purposes of example and that should not be taken as limiting the invention. Therefore, it is to be - - understood that, within the scope of the appended claims, the invention may be practiced other than as specifically described herein.

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