WO2018152065A1

WO2018152065A1 - Mechanically anisotropic tissue graft and 3d printing of a mechanically anisotropic tissue graft

Info

Publication number: WO2018152065A1
Application number: PCT/US2018/017837
Authority: WO
Inventors: Nicole L. BLACK; Jennifer A. Lewis
Original assignee: Harvard University
Current assignee: Harvard University
Priority date: 2017-02-14
Filing date: 2018-02-12
Publication date: 2018-08-23
Anticipated expiration: 2019-08-14

Abstract

A method of 3D printing a mechanically anisotropic tissue graft comprises flowing a polymeric ink formulation into a deposition nozzle, where the polymeric ink formulation includes (a) a block copolymer with hydrogen bonding segments and (b) a volatile solvent. A continuous filament comprising the polymeric ink composition is extruded from the deposition nozzle, and the volatile solvent evaporates rapidly upon extrusion. The continuous filament is deposited on a substrate as a deposited filament comprising the block copolymer while the deposition nozzle moves relative to the substrate in a print direction. Due to the rapid evaporation, at least a surface region of the deposited filament is substantially absent the volatile solvent. The deposition and/or extrusion induces elongation of the block copolymer and densification of the hydrogen bonding segments, and thus the deposited filament exhibits a higher average stiffness along the print direction than along a direction transverse to the print direction. Upon completion of the deposition, one or more of the deposited filaments are arranged in a predetermined architecture on the substrate, thereby defining a 3D printed mechanically anisotropic tissue graft.

Description

MECHANICALLY ANISOTROPIC TISSUE GRAFT AND 3D PRINTING OF A MECHANICALLY ANISOTROPIC TISSUE GRAFT

TECH N ICAL FI ELD

[0001] The present disclosure is related generally to three-dimensional (3D) printing and more particularly to 3D printing of mechanically anisotropic tissue grafts.

RELATED APPLICATIONS

[0002] The present patent document claims the benefit of priority under 35 U.S.C. 1 19(e) to U.S. Provisional Patent Application No. 62/458,818, filed on February 14, 2017, which is hereby incorporated by reference in its entirety.

FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

[0003] The invention described in this disclosure was made with

government support under Grant No. 2T32DC000038-26 awarded by the National Institutes of Health. The government has certain rights in the invention.

BACKGROU ND

[0004] The human body contains a wide variety of structurally and mechanically anisotropic soft tissues whose functions rely on this anisotropy. For example, the tympanic membrane (eardrum) contains radially- and circumferentially-directed collagen fibers. Studies have shown that radial stiffness plays a role in allowing the tympanic membrane to conduct sounds at both low and high frequencies. In addition to collagen fiber alignment, the alignment and directional elongation of cells themselves may have a

significant impact on the function of soft tissues. For example, in muscle tissue, cells are aligned for effective communication with each other and for proper mechanical functioning. In arteries, smooth muscle cells and vascular endothelial cells are oriented perpendicular to the direction of flow, like the collagen fibers in the artery wall. This orientation is critical for guiding the direction of blood flow. [0005] Despite the complex nature of many of the soft tissues in the human body due to their anisotropic cellular and extracellular matrix structures, current methods for cell alignment in tissue engineering have been limited. Many of these methods create architectures that are limited to one or two dimensions or directions of alignment. Additionally, the ability to align cells in vivo is difficult for many existing methods and generally requires a grafting material to guide the cells toward the preferred direction(s) of alignment in the tissue. However, many current graft materials lack the structure or properties needed to induce cellular alignment and instead result in isotropic remodeled tissues.

BRIEF SUMMARY

[0006] A mechanically anisotropic tissue graft and a method of 3D printing a mechanically anisotropic tissue graft are described in this disclosure.

[0007] The method entails flowing a polymeric ink formulation into a deposition nozzle, where the polymeric ink formulation includes (a) a block copolymer with hydrogen bonding segments and (b) a volatile solvent. A continuous filament comprising the polymeric ink composition is extruded from the deposition nozzle, and the volatile solvent evaporates rapidly upon extrusion. The continuous filament is deposited on a substrate as a deposited filament comprising the block copolymer while the deposition nozzle moves relative to the substrate in a print direction. Due to the rapid evaporation, at least a surface region of the deposited filament is substantially absent the volatile solvent. The deposition and/or extrusion induces elongation of the block copolymer and densification of the hydrogen bonding segments, and thus the deposited filament exhibits a higher average stiffness along the print direction than along a direction transverse to the print direction. Upon completion of the deposition, one or more of the deposited filaments are arranged in a predetermined architecture on the substrate, thereby defining a 3D printed mechanically anisotropic tissue graft. [0008] A mechanically anisotropic tissue graft comprises one or more deposited filaments arranged in a 2D or 3D architecture, where each of the deposited filaments comprises a block copolymer with densified hydrogen bonding segments. The densified hydrogen bonding segments are aligned along a longitudinal axis of the respective deposited filament. Each of the deposited filaments has higher average stiffness along the longitudinal axis thereof than along a direction transverse to the longitudinal axis, and thus the one or more deposited filaments are configured to induce cell elongation along a path of the 2D or 3D architecture.

BRI EF DESCRI PTION OF TH E DRAWINGS

[0009] FIG. 1 is a schematic of an exemplary 3D printing process.

[0010] FIG. 2A illustrates 3D printing of a polymeric ink formulation that comprises (a) a block copolymer with hydrogen bonding segments and (b) a volatile solvent; the inset figures illustrate the densification of the hydrogen bonding (or hard) segments that occurs during filament deposition.

[0011] FIG. 2B shows a larger-scale view of the 3D printing process of FIG.

2A, where one or more deposited filaments are arranged in a predetermined architecture to form a mechanically anisotropic tissue graft.

[0012] FIG. 3 is a schematic of an exemplary reaction of hard and soft segments and a chain extender to form a block copolymer with hydrogen bonding segments; in this example the block copolymer is a polyurethane.

[0013] FIG. 4 shows viscosity versus shear rate for three exemplary polymeric ink formulations to demonstrate the shear-thinning behavior of the ink formulations.

[0014] FIGs. 5A-5C show a 3D printed tympanic membrane tissue graft before (FIGs. 5A-5B) and after (FIG. 5C) seeding with cells; the tissue graft is printed from a biodegradable polymeric ink formulation comprising poly(ester urethane urea) (PEUU) in acetone.

[0015] FIGs. 6A and 6B show cell culture and biodegradability data comparing the block copolymer PEUU to polycaprolactone (PCL); specifically, FIG. 6A shows fibroblast proliferation as determined by MTS tetrazolium assay with cell normalization (n=3), and FIG. 6B shows degradation by hydrolysis measured in a phosphate buffered saline (PBS) bath at 37°C.

[0016] FIG. 7 shows stress-strain curves and stiffness values for tensile specimens prepared by 3D printing and strained either along the print path or transverse to the print path.

[0017] FIGs. 8A-8C show a comparison of GFP-HNDF (GFP-expressing human neonatal dermal fibroblast) alignment on a control PCL material as compared to PEUU on a cast sheet, a circular print path, and a parallel line print path. Cellular alignment is apparent on each of the printed PEUU grafts along the print path, but not on the PCL grafts.

[0018] FIG. 9A shows the print path (parallel lines) for 3D printing of polymeric ink formulations composed of PEUU mixed at 30 wt.% in solvents of different vapor pressures, specifically, acetone (FIG. 9B), HFIP (FIG. 9C), and ethyl acetate (FIG. 9D).

[0019] FIG. 9B shows an image of a graft 3D printed from PEUU/acetone after 7 days in culture with GFP-HNDFs.

[0020] FIG. 9C shows an image of a graft 3D printed from PEUU/HFIP after 7 days in culture with GFP-HNDFs.

[0021] FIG. 9D shows an image of a graft 3D printed from PEUU/ethyl acetate after 7 days in culture with GFP-HNDFs.

[0022] FIGs. 10A and 10B show PEUU/acetone ink formulations 3D printed in straight lines to form a tissue graft.

[0023] FIG. 10C shows the tissue graft of FIGs. 10A and 10B after seeding with cells on the underside of the tissue graft.

[0024] FIGs. 10D and 10E show PEUU/acetone ink formulations 3D printed in a circular/radial architecture to form a tissue graft.

[0025] FIG. 10F shows the 3D printed graft of FIGs. 10D and 10E after seeding with cells on the underside of the tissue graft. [0026] FIGs. 1 1 A and 1 1 B are images of a tissue graft printed in the shape of an "H" and seeded with HNDF cells; the images reveal that the cells tend to align along the print path of the graft.

[0027] FIG. 12A shows the structure of the temporomandibular joint (TMJ) cartilage disc.

[0028] FIG. 12B shows a schematic of the approximate collagen structure of the TMJ disc.

[0029] FIG. 12C shows a TMJ graft 3D printed from a PEUU/acetone ink formulation; the graft shows increased stiffness in the anterior-posterior (AP) direction.

[0030] FIG. 12D shows Live/Dead confocal images of hMSCs differentiated into chondrocytes on the 3D printed TMJ graft of FIG. 12C.

DETAILED DESCRIPTION

[0031] Block copolymers that have hydrogen bonding segments can exhibit "densification hardening" or "strain hardening" under mechanical deformation, which leads to an increased stiffness along the direction of strain. The inventors have discovered that 3D printing of an ink formulation comprising a strain-hardenable block copolymer and a rapidly evaporating solvent can provide a means to manipulate the stiffness anisotropy of printed filaments, thereby enabling the fabrication of anisotropic tissue grafts or scaffolds that promote elongation of cells along a desired direction (e.g., the path of filament deposition). Block copolymers may further have degradation rates and biocompatibility well suited for use in tissue engineering applications.

[0032] FIG. 1 A shows a close-up schematic of an exemplary 3D printing process, which may also be referred to as direct ink writing. Ink formulations that are suitable for 3D printing can be readily extruded through a deposition nozzle 102 moving relative to a substrate 104 to form a continuous filament 106 that substantially maintains its shape once deposited. As discussed below, suitable ink formulations exhibit shear-thinning behavior. During printing, the deposition nozzle 102 can be moved at a constant or variable speed along a desired print direction 114 while the substrate 104 remains stationary. Alternatively, the substrate 104 may be moved while the deposition nozzle 102 remains stationary, or both the deposition nozzle 102 and the substrate 104 may be moved.

[0033] In the new 3D printing process described in this disclosure, rapid evaporation of the solvent post-extrusion promotes the formation of deposited filaments with directional mechanical properties, which in turn facilitates the construction of anisotropic tissue grafts. Referring to FIGs. 2A and 2B, the method entails flowing a polymeric ink formulation 110 comprising a block copolymer with hydrogen bonding segments and a volatile solvent into a deposition nozzle 102. A continuous filament 106 comprising the polymeric ink composition 110 is extruded from the deposition nozzle 102, and the volatile solvent evaporates rapidly. The continuous filament 106 is deposited on the substrate 104 as a deposited filament 108 comprising the block copolymer while the deposition nozzle 102 moves relative to the substrate 104 in a print direction 114. The extrusion of the continuous filament 106 continues during deposition. Because the volatile solvent evaporates rapidly from the

polymeric ink formulation 110 as the continuous filament 106 is extruded, at least a surface region of the deposited filament 108 is substantially absent the volatile solvent.

[0034] The extrusion and/or deposition induces elongation of the block copolymer and densification (or strain hardening) of the hydrogen bonding segments, as shown schematically in FIG. 2A; thus, the deposited filament 108 exhibits an increased stiffness along the print direction 114. For example, the stiffness of the deposited filament 108 along the print direction 114 may be at least about 20% higher, or at least about 40% higher, than the stiffness of the deposited filament 108 transverse to the print direction 114. In one specific example, the deposited filament 108 may exhibit an average stiffness of at least about 60 MPa along the print direction 114. In view of the

directional stiffness, the deposited filament 108 may be said to exhibit stiffness anisotropy. Rapid evaporation of the solvent is essential for manipulation of the stiffness and other mechanical properties of the deposited filament 108 since strain hardening may not occur while the block copolymer is solvated. Since evaporation occurs more rapidly from the surface of the continuous filament 106 than from the interior, it is possible that strain hardening is enhanced within the surface region of the deposited filament 108.

[0035] At the conclusion of the 3D printing process, the deposited filament 108 with directional mechanical properties is arranged in a predetermined architecture 1 12 that defines a mechanically anisotropic tissue graft or scaffold 1 16, as shown for example in FIG. 2B. Given the stiffness and rigidity of the deposited filament 108, complex 3D patterns and high aspect ratio structures may be printed layer by layer. Various 2D and 3D architectures 1 12 can be printed to form mechanically anisotropic tissue grafts 1 16, as further described below. It is preferred but not necessary for the block copolymer of the deposited filament(s) 108 to be biodegradeable.

[0036] Block copolymers suitable for use in the polymeric ink formulation include hydrogen bonding segments or "hard segments" that can crystallize with each other through hydrogen bonding or other another weaker bonding interaction (e.g., pi stacking, ionic bonding, polar bonding, van der Waals bonding, or disulfide bonding). Polymers can hydrogen-bond between their chains if they contain both a hydrogen donor and a hydrogen acceptor.

Amide bonds, such as urethane, urea, and nylon, are the most common source of hydrogen bonding in block copolymers because these bonds usually contain both a hydrogen acceptor through the oxygen and a hydrogen donor through the nitrogen in the secondary amine. Block copolymers that include hydrogen bonding segments and which may be suitable for the polymeric ink formulation include polyurethane, nylon, polyurea, polyparaphenylene terephthalamide (Kevlar), cellulose, and proteins.

[0037] When the block copolymer is strained, the polymer chains tend to reorganize in the most energetically feasible way possible. This usually means that weaker interactions or links between polymer chains break and reform to create as many interactions or links as possible, a process referred to as densification. Densification during straining leads to an increased stiffness along the strain direction and may be referred to as densification hardening or strain hardening. In polyurethanes, urethane hard segments tend to density transverse to the strain direction. This makes the polymer stiffer in the strain direction, as more covalent bonds are oriented with the strain direction, and less stiff in the transverse direction, which is connected mainly by weaker hydrogen bonds.

[0038] Polyurethanes may be formed by reacting a hydrogen bonding segment or hard segment comprising a diisocyanate with a "soft segment" comprising a polyol. Diisocyanates tend to hydrogen-bond and crystallize with other diisocyanates in nearby polymer chains, giving polyurethane its elastomer-like properties. A chain extender may be included in the reaction to link polyurethane chains together to form a high molecular weight polymer. An exemplary two-step reaction of soft and hard segments and a chain extender is shown schematically in FIG. 3 and described in detail in the Examples below. Referring to FIG. 3, IPDI represents isophorene diisocyanate, a hard segment; PCL represents polycaprolactone diol, a soft segment; PU

represents putrescine, a chain extender; the sunbursts represent urethane bonds; and the crosses represent urea bonds. A suitable molar ratio of components is 1 :2:1 (soft segment: hard segment: chain extender). Other possible molar ratios may be selected from the following, where the chain extender is optional: 0.5-2 : 1 -5 : 0-2.

[0039] The diisocyanate may be selected from among isophorene diisocyanate (IPDI), methyl diphenyl diisocyanate (MDI), l-lysine diisocyanate (LDI), 1 ,4-butane diisocyanate (BDI), hexamethylene diisocyanate (HDI), and trimethylhexamethylene diisocyanate (TMDI). The chain extender may combine individual polyurethane chains through urea bonds or additional urethane bonds. The chain extender may comprise a diol or a diamine, such as ethylene glycol, 1 ,4-butanediol, 1 ,4-cyclohexanedimethanol, 1 ,2- ethanediamine, 1 ,4-butanediamine, combinations including 2-amino-1 - butanol, or another degradable linkage such as 2-hydroxyethyl-2- hydroxyproponoate.

[0040] The polyol may contain hydrolysable bonds, such as esters, ethers, or carbonates, which promote biodegradability. It has been demonstrated that urethane bonds may also be degraded and this degradation may be increased by exposure to protease, urease, esterase, and lipase enzymes that are produced by various cells in the body. Exemplary biodegradable

polyurethanes may include a diol (polyol) formed from polycaprolactone (PCL), poly(ethylene glycol) (PEG), poly(hexamethylene carbonate) (PHC), poly(ethylene oxide) (PEO), poly(propylene oxide) (PPO), polylactide, (PLA), polyglycolide (PGA), poly(hydroxybutyrate) (P3HB and P4HB), or amino acids.

[0041] The viscosity and flow properties of the polymeric ink formulation may be influenced by the molecular weight of the block copolymer as well as the amount or concentration of the block copolymer dissolved in the volatile solvent. A wide range of molecular weights may be suitable for 3D printing. Longer polymer chains may increase viscosity, which may promote solvent evaporation during extrusion, but it is also possible to utilize shorter chains and lower amounts of solvent. In general, the molecular weight of the block copolymer may range from about 1 kDa to about 100 kDa, where the molecular weight is more typically in the range from about 25 kDa to about 100 kDa after chain extension and from about 1 kDa to about 3 kDa before chain extension. Typically, the concentration of the block copolymer in the polymeric ink formulation is at least about 20 wt.% and may be as high as about 60 wt.%, depending on the molecular weight (chain length) of the block copolymer. Concentrations of at least about 30 wt.% or at least about 40 wt.% are typical. Given the block copolymer concentration, the volatile solvent may account for from about 40 wt.% to about 80 wt.% of the polymeric ink formulation.

[0042] Suitable volatile solvents evaporate rapidly (relative to water) and are capable of solubilizing the block copolymer. Solvent evaporation occurs nearly instantaneously from the surface of the continuous filament upon extrusion, where "nearly instantaneously" refers to a time scale typically in a range from about 0.01 second to about 1 second. Rapid solvent evaporation ensures that at least the surface region of the deposited filament is

substantially solvent-free, i.e., substantially absent the volatile solvent, such that the deposited filament that may undergo elongation and densification (and consequently anisotropic stiffening) during deposition. The extent of the surface region may be understood to be at least about 1 % and less than about 50% of the deposited filament in terms of volume, where the remaining volume of the filament (from 50 vol.% to less than about 99 vol.%) is considered to be the interior. Evaporation may also occur rapidly from the interior of the filament, depending on the diameter of the filament, the volatile solvent used, etc. The phrase "substantially absent the volatile solvent" means that no more than about 5 wt.% volatile solvent remains (e.g., in the surface region of the deposited filament) due to the rapid evaporation that occurs upon extrusion. In other words, at least the surface region of (and possibly an entirety of) the deposited filament includes the volatile solvent at a

concentration ranging from 0 wt.% (no solvent) up to about 5 wt.%. It is understood that this rapid evaporation may occur in ambient conditions (e.g., atmospheric pressure and room temperature (15-25°C)), without requiring exposure to heat. The volatile solvent employed in the polymeric ink formulation may be an organic solvent. For example, the volatile solvent may be selected from among acetone, toluene, hexafluoroisopropanol, ethyl acetate, diethyl ether, hexane, isopropanol, methylene chloride, ethanol, and methyl ethyl ketone. Generally speaking, the term "volatile solvent" as used herein may refer to an organic solvent that evaporates more rapidly than water.

[0043] To facilitate 3D printing, the polymeric ink formulation comprises a strain-rate dependent viscosity, and thus may be described as viscoelastic. More specifically, the polymeric ink formulation may be shear-thinning, a characteristic that provides a low viscosity at high shear rates (e.g., while passing through the deposition nozzle) and a higher viscosity at low shear rates (e.g., when deposited). The flow properties of exemplary polymeric ink formulations are shown in FIG. 4. Rapid evaporation of the volatile solvent further ensures an increase in viscosity and rigidity of the deposited filament.

[0044] Besides having inducible anisotropy, block copolymers with hydrogen bonding segments may also exhibit desirable mechanical properties for soft tissue engineering applications. Current tissue grafts are typically fabricated from hard, brittle polymers such as poly(capro-lactone) (PCL) and poly(lactic acid) (PLA), which have Young's moduli of 300-5,000 MPa, or from soft, ductile hydrogels such as gelatin or fibrinogen, which have Young's moduli of 1 -100 kPa. However, the elastic modulus of most collagenous tissues in the human body is between the range of these commonly used materials, at around 1-1 ,000 MPa. This may be problematic, as an ideal tissue graft should be able to integrate well into the mechanics of the surrounding tissue to prevent stress shielding if it is too stiff or tearing apart if it is too soft. Additionally, grafts that are too stiff may have difficulty being bent,

manipulated, and placed into small lesions, while grafts that are too soft may have difficulty being sutured into place, and thin grafts may lose their shape as they are being inserted. The polyurethanes and other block copolymers described in this disclosure have Young's modulus values (stiffnesses) much closer to the range of human soft tissues. For example, thermoplastic polyurethanes exhibit Young's modulus values between about 10 MPa and about 1000 MPa. Thus, the 3D printed tissue grafts may readily support the surrounding tissues as they are being remodeled and also be easily

manipulated and implanted into small cavities.

[0045] Thus, referring to FIGs. 5A and 5B, a 3D printed mechanically anisotropic tissue graft 1 16 comprises one or more deposited filaments 108 arranged in a 2D or 3D architecture or pattern 1 12, where each of the deposited filaments 108 comprises a block copolymer with densified hydrogen bonding segments. (The graft 1 16 shown in the photographic image of FIG. 5A is also shown schematically in FIG. 2B, with reference numerals). In the example of FIGs. 5A-5C, the tissue graft 116 is printed from a biodegradable polymeric ink formulation comprising poly(ester urethane urea) (PEUU) in acetone. Due to the 3D printing process, the densified hydrogen bonding segments are aligned along a longitudinal axis of the respective deposited filament 108. Thus, the deposited filaments 108 exhibit stiffness anisotropy, with each deposited filament 108 having a higher average stiffness along the longitudinal axis thereof than along a direction transverse to the longitudinal axis. For example, the average stiffness along the longitudinal axis may be at least about 20% higher, or at least about 40% higher, than the average stiffness along the direction transverse to the longitudinal axis. Consequently, the one or more deposited filaments 108 are configured to induce cell elongation along a path of the 2D or 3D architecture 112, which coincides with the print path during 3D printing. The mechanically anisotropic tissue graft 116 may be a tympanic membrane graft, as shown in FIGs. 5A-5C, or an articular cartilage graft, a muscle tissue graft, a nerve tissue graft, a vascular graft, or another type of tissue graft.

[0046] The block copolymer with "densified" hydrogen bonding segments that makes up the mechanically anisotropic tissue graft may have any of the characteristics described above for the block copolymer with hydrogen bonding segments. For example, the block copolymer may be biodegradable. In a preferred embodiment, the block copolymer may comprise polyurethane and the densified hydrogen bonding segments may comprise a diisocyanate, such as isophorene diisocyanate (IPDI), methyl diphenyl diisocyanate (MDI), I- lysine diisocyanate (LDI), 1 ,4-butane diisocyanate (BDI), hexamethylene diisocyanate (HDI), or trimethylhexamethylene diisocyanate (TMDI). The block copolymer may further include a soft segment comprising a diol formed from polycaprolactone (PCL), poly(ethylene glycol) (PEG), poly(hexamethylene carbonate) (PHC), poly(ethylene oxide) (PEO), poly(propylene oxide) (PPO), polylactide, (PLA), polyglycolide (PGA), poly(hydroxybutyrate) (P3HB and P4HB), or amino acids, and a chain extender comprising a diol or a diamine. [0047] After 3D printing, the mechanically anisotropic tissue graft may be seeded with cells for in vitro or in vivo remodeling. The cells selected for seeding depend on the architecture and intended function of the anisotropic tissue graft and may include neurons, fibroblasts, endothelial cells,

mesenchymal stem cells, smooth muscle cells, cardiac muscle cells, and/or skeletal muscle cells. Prior to seeding, it may be beneficial to plasma treat the anisotropic tissue graft using methods known in the art to kill bacteria and/or enhance hydrophilicity. Due to the stiffness anisotropy of the tissue graft imparted by 3D printing, some or all of the cells deposited on the graft may become elongated along the print direction (or along the longitudinal axis of the deposited filament ("filament axis"), which coincides with the print direction). Referring again to FIGs. 5A-5C, which show a 3D printed tympanic membrane graft before and after seeding with cells, the cells are elongated along the radial and circumferential directions of the graft.

[0048] The alignment of cells may be crucial for the alignment of fibrillar extracellular matrix proteins. For fibroblasts, cell elongation and spreading direction impacts the direction of collagen deposition. Atomic force microscopy (AFM) imaging of a 3D printed graft (prepared from a PEUU/acetone ink formulation) after 3 months of cellular remodeling reveals bands of what appear to be collagen fibers along the print path (or filament axis). This suggests that controlling cell alignment along the print path in turn affects collagen deposition and eventual collagen fiber organization also.

Examples

Synthesis of Biodegradable Polyurethane

[0049] Synthesis of an exemplary biodegradable polyurethane is done through a two-step melt reaction to create poly(ester urethane urea) (PEUU), where the soft segment is polycaprolactone diol (PCL) with hydrolysable ester bonds, the hard segment is isophorene diisocyanate (IPDI), and the chain extender is 1 ,4-diaminobutane (putrescine, PU), as shown schematically in FIG. 3. They are combined in a molar ratio of 1 :2:1 soft segment : hard segment : chain extender. First, the PCL (M_n=2000, Polysciences) is dried under vacuum at 50°C to remove the residual water before synthesis. Then, it is added to a 3-neck flask under the flow of nitrogen at 70°C to completely melt the polyol. Then, IPDI (TCI America) is added to the flask along with a few drops of stannous octoate catalyst (Spectrum Chemical). This first step forms the urethane bonds in the polymer. After 1 hour, the putrescine (Sigma- Aldrich) is mixed in dimethyl sulfoxide (DMSO) solvent at 20 wt.%, then added to the solution. The reaction continues for two more hours under nitrogen. This extends the chains with urea bonds, making the polymer stiffer and more biocompatible, as the urea bonds resemble peptide bonds to cells. After the reaction is complete, the reaction product is purified in water and ethyl acetate to remove residual monomers and dried in a vacuum oven at 70°C to obtain the final polymer, which in this example is PEUU. As shown in FIGs. 6A and 6B, this polymer demonstrates good biocompatibility with human neonatal dermal fibroblasts compared to a PCL control and a reasonable degradation rate, faster than that of PCL.

[0050] The soft segment, hard segment, and chain extender employed to form the biodegradable polyurethane can be varied as described above.

Other polymers that may be formed (e.g., by varying the soft segments) are poly(ester carbonate urethane urea) (PECUU), using PHC, and poly(ester ether urethane urea) (PEEUU), using PEG. The length of the chains from each synthesis reaction following the same protocol described above for the the original PEUU polymer, but substituting half of the PCL soft segment, is determined by gel permeation chromatography (GPC), as shown in Table 1 .

. Exemplary Biodegradable Polyurethanes

Polydispersity

[0051] Besides varying the components, the reaction can be modified to be done under argon or vacuum and for variations in temperature and time scale for each of the two steps. The entire reaction can also be conducted under a solvent such as dimethyl sulfoxide (DMSO) in lieu of the first step being performed as a melt reaction.

Creating a 3D Printable Ink with a Rapidly-Evaporating Solvent

[0052] The block copolymer may be combined with a volatile solvent to prepare the polymeric ink formulation. In these studies, acetone was used as the volatile solvent due to its biocompatibility, but other rapidly evaporating solvents, such as toluene, hexafluoroisopropanol, and methyl ethyl ketone, may also or alternatively be used. The block copolymer is generally combined in a Thinky cup with the volatile solvent at concentration between 20 wt.% and 60 wt.%, depending on its chain length and properties. To speed up the solvation of the polymer into the solvent, the two components can be heated below the solvent's boiling point (56°C for acetone) or mixed in the Thinky planetary centrifugal mixer. Together, these components give a shear-thinning ink as shown for example in FIG. 4, which is ideal for filamentary extrusion 3D printing. Rheology measurements were performed on a Discovery Series Hybrid Rheometer-3 (Texas Instruments) using a 40 mm flat cone geometry.

3D Printing of Block Copolymer Ink Formulation

[0053] Once the ink is formulated, it can be loaded into a syringe for filamentary extrusion 3D printing. The printer used in this investigation is a custom multimaterial printer (Aerotech), but others can also be used. A straight (or tapered) nozzle is connected to one side of the syringe (Nordson). Typically, nozzles with inner diameters (openings) from about 5 microns to about 200 microns are used to allow for faster evaporation of the solvent from the polymer because of a higher surface to volume ratio. The polymeric ink formulation is extruded via pneumatics onto a glass slide or a slide coated with a nonstick polymer, with pressures typically ranging from 5-100 psi depending on the block copolymer, solvent, and wt.% of the block copolymer in the ink formulation. Aerobasic G-code is used to program the print path of the nozzle in three dimensions. As the polymeric ink formulation is extruded, the solvent rapidly evaporates, leaving deposited filaments comprising just the block copolymer. Thus, as the nozzle translates, it pulls the deposited filament with it, densifying the hard segments (domains) of the block copolymer. After the filament(s) are printed to fabricate a tissue graft, the graft can be removed from the glass side and plasma treated (Diener ATTO) for 60 seconds on each side to make it more hydrophilic prior to cell seeding. The plasma treatment also works to clean the samples by killing bacteria on the surface.

Mechanical Properties of 3D-Printed PEUU

[0054] The mechanical properties of the biodegradable PEUU are determined by tensile tests of tensile specimens 3D printed from a polymeric ink formulation containing 30 wt.% PEUU in acetone using a 330-micron diameter nozzle. The cross-sectional areas are measured with micrometers, and then the specimens, shown schematically in FIG. 7, are subjected to tensile testing (Instron). As can be observed, the tensile specimen having a print path along the vertical direction, or along the longitudinal axis of the tensile specimen, yields the data along the upper stress-strain curves, and the tensile specimen having a print path along the horizontal direction, or transverse to the longitudinal axis of the tensile specimen, yields the data along the lower stress-strain curves. Young's moduli values (stiffness values) are calculated for each specimen from the stress-strain curves. The results show approximately a 179% increase in stiffness for the PEUU specimen having a print path aligned with the longitudinal axis instead of transverse to (e.g., normal to) the longitudinal axis. In the case of PLA control samples printed with the same vertical and horizontal print paths, the increase is only 119%. Importance of Choice of Polymer for Ink Formulation

[0055] As discussed above, use of a block copolymer with hydrogen bonding segments in the polymeric ink formulation promotes strain hardening during filament deposition and cell elongation upon seeding. The importance of the choice of polymer is investigated by printing ink formulations comprising polycaprolactone (PCL), which does not include hydrogen bonding segments, in a rapidly-evaporating solvent, and comparing the results to grafts prepared from PEUU in a rapidly-evaporating solvent.

[0056] Grafts are cast from both the PEUU and PCL ink formulations as well as printed in circular and parallel line print paths, as shown in FIGs. 8A- 8C. After 5 days in culture, GFP-HNDFs showed random spreading for both cast sheets (FIG. 8A). They also spread randomly on the confluent 3D printed PCL grafts independent of the print path (filament axis), suggesting that 3D printing of any polymer alone in a rapidly evaporating solvent may be insufficient to generate the mechanical anisotropy required to obtain cell elongation along a preferred direction. For both the circular and linear architectures (FIGs. 8B and 8C) of the printed PEUU grafts, the GFP-HNDFs oriented strongly along the direction of the print path/filament axis.

Importance of Rapid Solvent Evaporation

[0057] The importance of rapid solvent evaporation for cell elongation is investigated by printing grafts from polymeric ink formulations composed of PEUU mixed at 30 wt.% in solvents of different vapor pressures, including acetone (vapor pressure of 26 kPa), hexafluoroisopropanol (vapor pressure of 16 kPa), and ethyl acetate (vapor pressure of 13 kPa). The ink formulations are extruded through 200 μηι-diameter nozzles to print predetermined architectures (e.g., deposited filaments in parallel lines as shown in FIG. 9A) that could form all or part of a tissue graft, and the deposited filaments are seeded with cells. The images shown in FIGs. 9B-9D were taken after 7 days in culture with GFP-HNDFs, where FIG. 9B shows the seeded graft printed from the acetone-containing ink formulation, FIG. 9C shows the seeded graft printed from the HFIP-containing ink formulation, and FIG. 9D shows the seeded graft printed from the ethyl acetate-containing ink formulation. Only the tissue grafts prepared from the acetone-containing ink formulation showed strong cell elongation along the print path. It appears that only acetone, which has the highest vapor pressure of the three solvents, evaporates rapidly enough following extrusion to permit straining and densification of the PEUU during filament deposition.

Independence of Cellular Alignment from Graft Morphology

[0058] To demonstrate the independence of the cellular alignment from morphological or architectural features of the tissue grafts, polymeric ink formulations comprising PEUU and acetone are 3D printed in straight lines (FIGs. 10A and 10B) and in a circular/radial architecture (FIGs. 10D and 10E). Cells are then seeded on the underside of the tissue grafts, which are fundamentally flat due to compression on the glass. Despite the absence of fiber curvature on the underside of the deposited filaments, the cells still aligned along the print path or filament axis of the bottom-most layer of the grafts, as can be seen in FIGs. 10C and 10F.

[0059] In a second experiment, PEUU/acetone inks are cast into a 3D printed PDMS mold to create an exaggerated structure of fibers complete with ridges but without the mechanical strain imparted to deposited filaments during 3D printing. Images reveal that cells do not align along the cast fibers, demonstrating that it is not the architecture of the 3D printed tissue grafts that induce cell alignment, but rather the microstructure (e.g., stiffness anisotropy) of the polymer itself. This experiment suggests that it is insufficient to cast or otherwise manufacture texture at this scale into grafts for cell alignment. The changes in microstructure (e.g., domain densification) and mechanical properties that occur during filament extrusion and/or deposition are critical.

Demonstrations of Cellular Alignment

[0060] Alignment of cells along the print paths of anisotropic tissue grafts of varying complexity has been demonstrated, including circular grafts, circular/radial grafts, and a letter "H," as shown in FIGs. 11 A and 11 B. In all shapes printed, the cells tend to align along the print path of the graft, including along curved lines. These results suggest great potential for 3D printing of biodegradable polyurethane (or other block copolymers with hydrogen bonding segments) in a rapidly evaporating solvent as a means of constructing simple or complex tissue grafts capable of inducing cell elongation along desired directions. As the polymer degrades, the graft may be replaced by a cellular and extracellular architecture that closely mimics the design of the print path.

(a) Tympanic Membrane Grafts

[0061] As mentioned above, radial stiffness in the tympanic membrane (TM) imparted by collagen fibers in the lamina propria is crucial to sound conduction at both high and low frequencies. Previous investigations have shown that TM grafts 3D printed from various polymers (e.g., PDMS, PLA, and PCL) which do not strain harden upon extrusion do not promote alignment of cells or anisotropic collagen I deposition. In contrast, TM grafts 3D printed from PEUU/acetone ink formulations in an architecture that includes circu inferential ly directed lines on the bottom of the graft and radially directed lines on top of the graft are able to induce cell alignment, as shown in FIGs. 5A-5C.

(b) Temporomandibular Joint (TMJ) Grafts

[0062] The temporomandibular joint (TMJ) contains an articular cartilage disc with a fibrous collagen structure that gives it anisotropic mechanical properties. One researcher measured this stiffness difference to be 76 MPa in the anterior-posterior (AP) direction and only 3 MPa in the medial-lateral (ML) direction, as can be understood in view of FIG. 12A. This disc is composed of mainly of fibroblasts, which make up about 70% of the cells, and chondrocytes which make up about 30% of the cells. To mimic the collagenous architecture of the disc, an Aerobasic G-code print path shown in FIG. 12B was designed with lines parallel to the AP direction in the center with a border of radial and circular fibers at the edges. The program allows for parameters, such as length, width, fiber spacing, and number of layers, to be rapidly varied and customized. The overall dimensions of the TMJ disc were taken from previous literature to be a width of 13 mm and length of 19 mm, with a thickness of 1-2 mm. An exemplary 3D printed TMJ disc is shown in FIG. 12C. The ability of human mesenchymal stem cells (hMSCs) differentiated into chondrocytes to proliferate and elongate on the PEUU grafts was examined. Preliminary Live/Dead imaging shows potential for the chondrocytes to grow on both PCL printed filaments and PEUU printed filaments. However, the cells elongate only on the PEUU filaments, as expected, and spread randomly on the PCL filaments. When seeded onto the TMJ grafts, the hMSCs appear to elongate along the print direction (longitudinal axis) for the different filament

orientations, as evidenced in FIG. 12D.

[0063] Although the present invention has been described in considerable detail with reference to certain embodiments thereof, other embodiments are possible without departing from the present invention. The spirit and scope of the appended claims should not be limited, therefore, to the description of the preferred embodiments contained herein. All embodiments that come within the meaning of the claims, either literally or by equivalence, are intended to be embraced therein.

[0064] Furthermore, the advantages described above are not necessarily the only advantages of the invention, and it is not necessarily expected that all of the described advantages will be achieved with every embodiment of the invention.

Claims

1. A method of 3D printing a mechanically anisotropic tissue graft, the method comprising:

flowing a polymeric ink formulation into a deposition nozzle, the polymeric ink formulation including a block copolymer with hydrogen bonding segments and a volatile solvent;

extruding a continuous filament comprising the polymeric ink

composition from the deposition nozzle, the volatile solvent evaporating rapidly upon extrusion; and

depositing the continuous filament on a substrate as a deposited filament comprising the block copolymer while the deposition nozzle moves relative to the substrate in a print direction, at least a surface region of the deposited filament being substantially absent the volatile solvent,

wherein the deposition and/or extrusion induces elongation of the block copolymer and densification of the hydrogen bonding segments, the deposited filament thereby exhibiting a higher average stiffness along the print direction than along a direction transverse to the print direction, and

wherein, upon completion of the deposition, one or more of the deposited filaments are arranged in a predetermined architecture on the substrate, thereby defining a 3D printed mechanically anisotropic tissue graft.

2. The method of claim 1 , wherein the block copolymer is

biodegradable.

3. The method of claim 1 or 2, wherein the block copolymer is selected from the group consisting of: polyurethane, nylon, polyurea, polyparaphenylene terephthalamide (Kevlar), cellulose, and proteins.

4. The method of any one of claims 1-3, wherein the block copolymer comprises polyurethane and the hydrogen bonding segments comprise a diisocyanate.

5. The method of any one of claims 1-4, wherein the block copolymer further comprises a soft segment comprising a diol formed from polycaprolactone (PCL), poly(ethylene glycol) (PEG), poly(hexamethylene carbonate) (PHC), poly(ethylene oxide) (PEO), poly(propylene oxide) (PPO), polylactide, (PLA), polyglycolide (PGA), poly(hydroxybutyrate) (P3HB and P4HB), or amino acids.

6. The method of any one of claims 1-5, wherein the block copolymer further comprises a chain extender comprising a diol or a diamine.

7. The method of any one of claims 1-6, wherein the volatile solvent is selected from the group consisting of acetone, toluene,

hexafluoroisopropanol, ethyl acetate, diethyl ether, hexane, isopropanol, methylene chloride, ethanol, and methyl ethyl ketone.

8. The method of any one of claims 1-7, wherein the block copolymer is present in the polymeric ink formulation at a concentration from about 20 wt.% to about 60 wt.%.

9. The method of any one of claims 1-8, wherein the polymeric ink formulation is shear thinning.

10. The method of any one of claims 1 -9, wherein the 3D printed mechanically anisotropic tissue graft is selected from the group consisting of: tympanic membrane graft, articular cartilage graft, muscle tissue graft, nerve tissue graft, and vascular graft.

11. The method of any one of claims 1-10, further comprising seeding the 3D printed mechanically anisotropic tissue graft with cells.

12. The method of claim 11 , wherein the cells are selected from the group consisting of: neurons, fibroblasts, endothelial cells, mesenchymal stem cells, smooth muscle cells, cardiac muscle cells, and skeletal muscle cells.

13. The method of claim 11 or 12, further comprising, prior to seeding, plasma treating the 3D printed mechanically anisotropic tissue graft to kill bacteria and/or enhance hydrophilicity.

14. The method of any one of claims 1-13, wherein the average stiffness of the deposited filament along the print direction is at least about 60 MPa.

15. The method of any one of claims 1-14, wherein the average stiffness of the deposited filament along the print direction is at least about 20% higher than the average stiffness of the deposited filament transverse to the print direction.

16. A mechanically anisotropic tissue graft comprising:

one or more deposited filaments arranged in a 2D or 3D architecture, each of the deposited filaments comprising a block copolymer with densified hydrogen bonding segments, the densified hydrogen bonding segments being aligned along a longitudinal axis of the respective deposited filament, and wherein each of the deposited filaments has a higher average stiffness along the longitudinal axis thereof than along a direction transverse to the longitudinal axis, the one or more deposited filaments thereby being

configured to induce cell elongation along a path of the 2D or 3D architecture.

17. The mechanically anisotropic tissue graft of claim 16, wherein the average stiffness along the longitudinal axis is at least about 20% higher than the average stiffness along the direction transverse to the longitudinal axis.

18. The mechanically anisotropic tissue graft of claim 16 or 17, further comprising cells on the one or more of the deposited filaments, at least some of the cells being elongated along the longitudinal axis of the respective deposited filament.

19. The mechanically anisotropic tissue graft of claim 18, wherein the cells are selected from the group consisting of: neurons, fibroblasts, endothelial cells, mesenchymal stem cells, smooth muscle cells, cardiac muscle cells, and skeletal muscle cells.

20. The mechanically anisotropic tissue graft of any one of claims 16-19, wherein the block copolymer is biodegradable.

21. The mechanically anisotropic tissue graft of any one of claims 16-20, wherein the block copolymer is selected from the group consisting of: polyurethane, nylon, polyurea, polyparaphenylene terephthalamide (Kevlar), cellulose, and proteins.

22. The mechanically anisotropic tissue graft of any one of claims 16-21 , wherein the block copolymer comprises polyurethane and the densified hydrogen bonding segments comprise a diisocyanate.

23. The mechanically anisotropic tissue graft of any one of claims 16-22 selected from the group consisting of: tympanic membrane graft, articular cartilage graft, muscle tissue graft, nerve tissue graft, and vascular graft.