Patent Application
20250041487
This invention relates to tissue engineering constructs and methods of making thereof.
The regeneration of large bone defects caused by skeletal injuries, diseases, or congenital disorders remains a significant clinical problem. Over 0.5 million and 2 million bone grafting procedures are done in the US and worldwide, respectively every year. Autologous bone is the gold standard for bone grafting. However, an additional surgery is needed to harvest the autologous bone from the donor site, the amount of harvested bone is limited for reconstruction of large defects and the donor site may become morbid. Allogenic grafts have been widely used as an alternative to autologous grafts for bone regeneration. However, the long-term failure rate of allogenic grafts in treatment of large critical bone defects is 25%˜60% due to various complications. In addition, the use of frozen allografts suffers from a potential risk of disease transmission according to the Center for Disease Control and Prevention (CDC). Commercially available demineralized bone matrix (DBM) contains osteo-inductive factors but DBM alone does not provide the structural and mechanical support for reconstruction of large bone defects. Therefore, synthetic scaffolds as bone graft substitutes have attracted attention in recent years.
An ideal scaffold for bone tissue engineering is biocompatible, bioresorbable, mechanically stable, porous, osteo-conductive and osteo-inductive. Biocompatible, bioresorbable, and FDA-cleared polyesters including polycaprolactone (PCL), polyglycolic acid (PGA), and polylactic acid (PLA) and their copolymers (e.g. PLGA) are the most widely used synthetic polymers in bone tissue engineering. Several methods such as molding, solvent casting/porogen leaching, gas foaming, laser drilling and 3D printing have been applied to make porous polyester-based scaffolds. Among these techniques, 3D printing offers a precise control over the architecture and porosity of the scaffold. A well-controlled porosity of 3D printed scaffolds is particularly important for bone tissue engineering, because the presence of interconnected pores with individual pore size larger than 300 μm is essential for cell migration and bone ingrowth.
The osteo-conductivity and mechanical properties of the polyester-based scaffolds have been augmented by incorporation of calcium phosphate bioceramics. For example, the osteogenic differentiation of mouse preosteoblast cells (MC3T3-E1) on 3D printed PCL/β-tricalcium phosphate (TCP) substrates was significantly higher than on pristine PCL substrates. The inventors, previously, showed that the Young's modulus of 3D printed PCL-TCP scaffolds was tunable in 12 to 188 MPa range by changing the TCP content and scaffold porosity. Further, the clinically available electron beam sterilization did not adversely affect the mechanical and bioactive properties of PCL-TCP scaffolds.
Although 3D printed polymer/ceramic scaffolds are biocompatible, bioresorbable, mechanically stable, porous and osteo-conductive, they lack the osteo-inductive factors to stimulate osteogenic differentiation and accelerate bone healing. Therefore, there is a need to incorporate osteo-inductive proteins into 3D printed scaffolds particularly for treatment of large bone defects. Surface coating has been used to immobilize proteins on the surface of 3D printed scaffolds for tissue engineering applications. However, the loading of proteins on thin coatings is typically limited and the release rate is fast. For instance, the loading of BSA on 3D printed hydroxyapatite-based scaffolds coated with chitosan and sodium hyaluronate by layer-by-layer (LBL) deposition, was lower and the release was faster than uncoated scaffolds.
Hydrogels have been used for protein delivery or cell encapsulation in tissue engineering applications. A 3D polymeric network of hydrogels with a large water content provides a platform for adequate loading and sustained release of proteins. However, inferior stiffness and structural integrity of hydrogels limit their use as stand-alone 3D scaffolds. In addition, incorporation of soft hydrogels into rigid 3D printed scaffolds is challenging due to mechanical property mismatch at the interface. Filling porous 3D printed scaffolds with hydrogel precursor solution followed by photo or thermal induced gelation has been used to incorporate hydrogels into rigid scaffolds. For example, a surface tension-assisted method was used to fill the pores of 3D printed constructs with photo-crosslinkable methacrylated gelatin hydrogel. Multi-material 3D printing has been used to manufacture porous polymer/hydrogel composite scaffolds.
Preservation of the porous structure of the scaffold after incorporation of hydrogel is essential for cell migration, tissue integration, and vascularization for diffusion of nutrients and oxygen in tissue engineering applications. An integrated tissue-organ printer was used to sequentially print a gelatin/fibrinogen-based hydrogel along with PCL structural support. Stanford University developed a 3D hybrid bioprinting technology (Hybprinter) and used it for printing composite scaffolds from PCL and polyethylene glycol diacrylate (PEGDA) hydrogel. Despite its technological significance, multi-material printing requires a long fabrication time due to multiple iterations between the materials during printing, and a specialized expensive 3D printer. Furthermore, concurrent printing of polymer/ceramic and hydrogels hinders the scaffold surface treatment and improving the integration between soft and rigid materials at the interface.
This invention in one example provides a porous or a non-porous biologics-loaded multimaterial construct, hereafter referred to as Hybrid Tissue Engineering Construct (HyTEC) for applications in regenerative medicine and treatment of diseases.
Constructs and devices made of polymers, ceramics, metals, or composites in porous or non-porous forms have been widely used as implants in regenerative medicine. A number of techniques, including 3D printing and casting, have been used to manufacture porous implants. Also, coating techniques including layer-by-layer coating or adhesive coating have been used to load biomolecules on the surface of porous implants. However, these coating techniques only allow loading of a limited amount of biomolecules. Loading a large or tunable dose of biomolecules on implants is particularly important since the effective dose of biomolecules is often high in-vivo and could be different for various indications.
Biologics (biomolecules, drugs and/or cells) could be loaded on implants via filling the porous structure of the implant with a biologics-loaded hydrogel. However, filling the porous space of implants with a hydrogel closes the pores and inhibits or mitigates cell recruitment and migration, vascular invasion, tissue regeneration, and integration with surrounding tissues.
To address at least this concern, the inventors of this invention have developed a strategy to engineer a HyTEC that enables incorporation of biologics through a uniform thick hydrogel layer onto porous scaffolds while retaining interconnected open pores, or onto non-porous implants (
The surface of porous or non-porous scaffolds is treated in three consecutive steps to (
A layer of hydrogel is loaded on the surface of scaffolds through a surface-initiated physical crosslinking followed by covalent crosslinking.
Sodium hydroxide (NaOH) treatment and freezing/thawing were used to increase the surface hydrophilicity/reactivity in PCL-TCP scaffolds. Other treatment methods including plasma or acid treatment could also be used to increase the surface hydrophilicity/reactivity/roughness.
For improving the hydrogel adhesion, the surface is coated with a molecule that has a covalently linkable functional group. For instance, reactive Aminopropyl methacrylamide (APMA), and Gelatin methacrylate (GelMA) have been conjugated to the surface of PCL-TCP scaffolds using carbodiimide chemistry (
To stimulate surface-initiated physical crosslinking, calcium chloride (CaCl2) or calcium sulfate (CaSO4) was deposited on the surface of the implants. Other salts of divalent cations (e.g. Ca2+, Mg2+, Sr2+) or multivalent cations (e.g. Ti4+ or Al3+) could also be used for surface initiated physical crosslinking.
After these three steps of surface treatment, the scaffolds are dipped into a hydrogel precursor solution containing alginate, covalently reactive macromonomers, an initiator, and biologics (biomolecules, drugs, and/or cells). Polyethylene glycol dimethacrylate (PEGDMA) and GelMA were used as covalently reactive macromonomers (
When the surface treated scaffolds are dipped into the hydrogel precursor solution, calcium ions diffuse from the surface to the solution, crosslink alginate at the proximity of the surface, and make a hydrogel layer on the surface. The macromonomers within the physically crosslinked hydrogel are then covalently crosslinked in the next step to form a stiff interpenetrating network. A chemical initiator (APS/TEMED) and a photoinitiator (Lithium phenyl-2,4,6-trimethylbenzoylphosphinate) have been used for making porous and non-porous HyTECs, respectively (
The hydrogel loading and hydrogel thickness are tuned by changing the process parameters. For instance, the thickness of hydrogel layer on porous PCL-TCP scaffolds was tuned by changing the NaOH surface treatment time and the CaCl2) concentration in the solution that was used for calcium deposition. The coating thickness within a construct could be tuned/controlled from zero to high value spatially by dipping different parts into different solutions.
The 3D printed porous PCL-TCP scaffolds with different porosities remained porous after hydrogel loading (
Depending on the application, the HyTEC can be cellular or cell free, with or without therapeutic biomolecules. It can be fresh, freeze or freeze-dried. The scaffold can also be porous or non-porous and be made from polymers (e.g. polyesters), metals, ceramics, or composites. The hydrogel macromonomer, gelation initiator, salt for calcium deposition, and the material used for surface treatment can be changed depending on the scaffold surface chemistry and application.
Embodiments of the invention could be applied or used in the following ways without any limitation to be scope of the invention. HyTEC used for delivery of therapeutics including cells and/or biomolecules along with a structural support and a defined geometry for applications in regenerative medicine. Examples are as follows:
Embodiments of the invention are advantageous over existing approaches and constructs. A large dose, a broader spectrum of dose, or a variety of therapeutics or biologics can be loaded on porous (or non-porous) constructs using HyTEC technology as opposed to methods that are based on thin coatings (e.g. layer-by-layer coating or coating the construct with an absorbent). In addition, cells can be encapsulated in HyTEC as opposed to constructs with thin coating.
The advantage of HyTEC technology over multimaterial printing is as follows. Despite its technological significance, multimaterial printing requires a long fabrication time due to multiple iterations between the materials during printing, a specialized expensive 3D printer, and limited selections of processing parameters due to the nature of various printing mechanisms. Furthermore, concurrent printing of polymer/ceramic and hydrogels hinders the scaffold surface treatment and improving the integration between soft and rigid materials at the interface.
In one embodiment, to slow down the release of therapeutics, the bioactive implants (e.g. HyTEC constructs) could be coated with a resorbable polyester (e.g. PCL, PLA, or PLGA) or other resorbable polymers (e.g. polyurethanes). HyTEC stands for hybrid tissue engineered construct, which is a bioactive implant. A schematic representation of the method that is used to coat the HyTEC constructs is shown in
In another example, the present invention provides a method of forming a tissue engineering construct. A scaffold with a surface and a surface area is provided. The surface of the scaffold is treating to increase the surface area of the scaffold. Optionally the surface area of the scaffold is prepared to facilitate a chemical cross-linking to the surface area by coating the surface area with covalently linkable molecules. The surface area of the scaffold is prepared to facilitate surface-initiated physical cross-linking by depositing a salt onto the surface area or the optionally coated surface area. A hydrogel precursor solution is provided/prepared containing charged polymers, covalently reactive macromonomers, an initiator and biologics. A physically cross-linked hydrophilic hydrogel network is formed onto the surface of the scaffold by immersing the prepared scaffold into a hydrogel precursor solution. The forming is controlled by a release of salt-ions from the surface area and physically cross-linking the charged polymers with the released salt-ions. During the formation the biologics becomes trapped and thereby hosted within the physically cross-linked hydrogel network. The scaffold with the physically cross-linked hydrophilic hydrogel network is removed from the hydrogel precursor solution. The covalently reactive macromonomers are chemically cross-linked within the physically cross-linked hydrophilic hydrogel network to strengthen the physically cross-linked hydrophilic hydrogel network itself and to the scaffold. Optionally (following the optional preparation of the surface area infra) the coated covalently linkable molecules are chemical cross-linked with the covalently reactive macromonomers to increase adhesion of the chemically and physically cross-linked hydrophilic hydrogel network to the scaffold.
In a further step, freeze or freeze-drying of the tissue engineering construct can be performed if needed.
In a further step, the surface area of the scaffold can be treated to increase the hydrophilicity and/or roughness of the surface of the scaffold.
The scaffold can be an interconnected porous structure and as such the methods steps can then be controlled for the hydrophilic hydrogel network to be physically and chemically crosslinked and chemically bound to the interconnected porous structure of the scaffold. The pores of the interconnected porous structure can then be preserved by these controlled method steps to allow the pores to also be house the biologics.
In still a further step, the tissue engineering construct can be coated with one or more coating layers.
In still a further step, the scaffold has an interconnected porous structure and where the coating controls pore size of the interconnected porous structure.
In another embodiment, a tissue engineering construct is provided. The construct distinguished a scaffold with a surface and a treated surface area for increased surface area. A hydrophilic hydrogel network is physically cross-linked via charged polymers and salt-ions onto the treated surface area. Biologics is trapped and thereby hosted within the physically cross-linked hydrogel network. Covalently reactive macromonomers are chemically cross-linked within the physically cross-linked hydrophilic hydrogel network to strengthen the physically cross-linked hydrophilic hydrogel network itself and to the scaffold.
In a variation of the construct, the surface area can be coated with covalently linkable molecules which are chemical cross-linked with the covalently reactive macromonomers to increase adhesion of the chemically and physically cross-linked hydrophilic hydrogel network to the scaffold. The construct can have one or more coatings. The scaffold can be an interconnected porous scaffold where the biologics are then also hosted with pores of the interconnected porous scaffold.
Fused deposition modeling is a powerful method for printing 3 dimensional (3D) bioresorbable scaffolds and medical devices with well-controlled porosity, internal microstructure, and overall geometry for biomedical applications. However, proteins and live cells are not able to withstand the hot extrusion. In a further characterization of the invention a hybrid tissue engineering construct (HyTEC) is engineered that enables incorporation of biologics (e.g. proteins and cells) through a uniform thick hydrogel layer onto 3D printed scaffolds while retaining interconnected open pores, or onto non-porous implants. A 3D printed biodegradable polycaprolactone-β-tricalcium phosphate (PCL-TCP) was used as a model porous scaffold, a PCL-TCP rod as a model non-porous implant, and bone morphogenetic protein-2 (BMP-2) as a model protein for bone tissue engineering application. The surface of PCL-TCP constructs was treated in three consecutive steps to increase hydrophilicity, improve hydrogel adhesion and stimulate surface-initiated crosslinking. A layer of hydrogel was loaded on the surface of scaffolds through a surface-initiated physical crosslinking followed by covalent crosslinking. The results showed that surface treatment did not adversely affect the mechanical and surface properties of the scaffolds but improved the adhesion of hydrogel to the surface. The average thickness of the loaded hydrogel layer was controlled in the range of 100-600 μm by adjusting surface treatment parameters. In addition, 3D printed scaffolds with 50-80% porosity remained porous after hydrogel loading with pore sizes ranged from 140 to 1100 μm. Cell viability and proliferation tests using two cell types (hMSCs and C3H10) showed that hydrogel loading did not adversely influence the biocompatibility of scaffolds. BMP-2-laden hydrogel loaded scaffolds released BMP-2 in a sustained manner over 35 days. Freeze-drying and E-beam sterilization of hydrogel loaded PCL-TCP scaffolds did not adversely affect the mechanical properties of scaffolds but negatively impacted the amount of released active BMP-2. The amount of released active BMP-2 from sterilized freeze-dried HyTEC constructs was improved by 2 folds by using a split dose E-beam strategy. A thick hydrogel layer enabled loading encapsulated live cells on HyTEC constructs with over 92% cell viability after 7 days. In summary, the HyTEC strategy introduced in this study holds great promises in porous (or non-porous) polyester-based 3D printed tissue engineering scaffolds with improved payload capacity of biological substances while maintaining interconnected open pores for improved tissue integration and engraftment.
If needed, for further interpretation of the gray-scale in the drawings the reader is referred to the priority document(s) for each of the respective figures.
The following detailed description is exemplary embodiments of the method of forming/making the tissue engineering construct and the structural features of the tissue engineering construct. In general, the following definitions of terms can be used as a guidance within the scope of the invention.
A facile method was developed for manufacturing PCL-TCP/hydrogel composite scaffolds, hereafter referred to as hybrid tissue engineering construct (HyTEC). After 3D printing, the surface of PCL-TCP scaffolds was treated in three steps to increase hydrophilicity, improve hydrogel adhesion and stimulate surface-initiated crosslinking. A physically crosslinked hydrogel was then loaded on the scaffolds followed by covalent crosslinking of the hydrogel to form a stable interpenetratable network. The effects of surface treatment, processing parameters and freeze-drying on hydrogel loading and preservation of porous structure of the scaffold after manufacturing were investigated. Further, the adhesion between the PCL-TCP and hydrogel as well as the release kinetics of BMP2 protein encapsulated in HyTEC were evaluated. Moreover, the biocompatibility and osteo-inductive potential of BMP2 loaded HyTEC were studied. The method/strategy could be used for incorporating a wide range of hydrogels into porous polyester-based scaffolds as well as coating non-porous polyester-based constructs. As an example, at the end of this description the inventors demonstrated the efficacy of the method/strategy for coating non-porous PCL-TCP rods by modifying hydrogel and crosslinking mechanism for sustained release of BMP2 protein and for cell encapsulation.
The following description of materials and methods are exemplary embodiments.
Medical-grade polycaprolactone (PCL, Mn=80 kDa) was purchased from Sigma-Aldrich. β-TCP nano-powder with average particle size of 100 nm (TCP) was received from Berkeley Advanced Materials Inc. Dimethylformamide (DMF), sodium hydroxide (NaOH) and ethanol were purchased from Fisher Scientific Inc. N-(3-Dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC), N-Hydroxysulfosuccinimide (NHS), 2-(N-Morpholino)ethanesulfonic acid (MES), N-(3-Aminopropyl)methacrylamide hydrochloride (APMA), Ammonium persulfate (APS), N,N,N′,N′-Tetramethylethylenediamine (TEMED), gelatin type A, Heparin, ninhydrin and Triton X-100 were purchased from Sigma-Aldrich. Ninhydrin reagent was prepared by dissolving 20 mg/mL ninhydrin in ethanol. Polyethylene dimethacrylate (PEGDMA, Mn=1000 gr/mol) was received from Polyscience, Inc. Sodium alginate (alginate, 500GM) was purchased from Pfaltz & Bauer Inc. Human BMP2 protein was provided by Medtronic. Calcium colorimetric assay (MAK022), CCK-8 kit and Human BMP2 ELISA kit were purchased from Sigma-Aldrich. QuantiChrom ALP Assay Kit was received from BioAssay Systems LLC. Quant-it PicoGreen assay kit was purchased from Thermo Fisher Scientific.
PCL-TCP filament with PCL to TCP weight ratio of 80:20 was synthesized as described by Bruyas (Bruyas et al., Effect of Electron Beam Sterilization on Three-Dimensional-Printed Polycaprolactone/Beta-Tricalcium Phosphate Scaffolds for Bone Tissue Engineering, Tissue Eng Pt A (2018)). Briefly, 80 gr of PCL and 20 gr of TCP were dissolved in 800 mL and 400 mL of DMF, respectively at 80° C. with continuous stirring for 3 hours. The PCL and TCP solutions were then mixed and stirred for one hour, followed by precipitation in 4 liters of water to make PCL-TCP composite. The PCL-TCP composite was rinsed with water to remove the residual solvent and air dried at ambient temperature for 24 hours. The dried PCL-TCP composite was cut into pellets and extruded using an in-house built screw extruder as described by Bruyas (Bruyas et al, Systematic characterization of 3D-printed PCL/beta-TCP scaffolds for biomedical devices and bone tissue engineering: Influence of composition and porosity, J Mater Res 33(14) (2018) 1948-1959). PCL-TCP scaffolds were 3D printed using a Lulzbot Mini (Aleph Objects Inc, USA) with a nozzle diameter of 500 μm. For surface characterization, non-porous disks with 10 mm diameter and 600 μm thickness and for all other tests porous cylinders with 10 mm diameter and 5 mm height were printed. The 3D models were designed using SolidWorks (SolidWorks Corp.) and sliced using Cura software. For printing 0%, 30%, 50%, 60%, 70% and 80% porous scaffolds, strut distances of 0.4, 0.53, 0.80, 1.00, 1.25 and 2.00 mm were used. The printing temperature, layer thickness and printing speed were set to 160° C., 200 μm and 5 mm/s, respectively as described by Bruyas (Bruyas et al., Effect of Electron Beam Sterilization on Three-Dimensional-Printed Polycaprolactone/Beta-Tricalcium Phosphate Scaffolds for Bone Tissue Engineering, Tissue Eng Pt A (2018). To synthesize methacrylated gelatin (GelMA) macromonomer, gelatin was dissolved in DI water (10% w/v) at 50° C. Methacrylic anhydride was added to gelatin solution at a molar ratio of 100:1 (methacrylic anhydride:gelatin) and the solution was allowed to react under stirring for 1 hr at 50° C. The mixture was then 5× diluted with DI water and dialyzed against DI water using a dialysis tube (Spectrum Laboratories, Rancho Dominquez, CA) with 6-8 kDa molecular weight cutoff for 3 days at 40° C. The GelMA solution was then freeze-dried and stored at −80° C.
To synthesize methacrylated heparin (HepMA), 1 gr heparin was dissolved in 100 mL MES buffer (100 mM). 5 mL MES buffer containing 45 mg EDC and 30 mg NHS was then added to the heparin solution to activate the carboxylic acid groups as described by Jeon (Jeon et al, Affinity-based growth factor delivery using biodegradable, photocrosslinked heparin-alginate hydrogels, J Control Release 154(3) (2011) 258-66). After 1 hr reaction at room temperature, 25 mg APMA in 1 mL MES was added to the solution and allowed to react for 2 hr at room temperature. The methacrylated heparin solution was then dialyzed against DI water using a dialysis tube (Spectrum Laboratories, Rancho Dominquez, CA) with 6-8 kDa molecular weight cutoff for 3 days at ambient temperature, lyophilized, and stored at −80° C.
The procedure for scaffold surface treatment and hydrogel formation is shown schematically in
The hydrogel precursor solution was prepared by dissolving PEGDMA (10-30% wt/vol), alginate (1.5% wt/vol) and 1 mg/mL APS in DI water. The surface treated scaffolds were dipped into the hydrogel precursor solution for 1 min at room temperature and then centrifuged at 1000 rpm for 1 min to remove the residual precursor solution. At this step, a layer of hydrogel was formed on the scaffold surface due to the diffusion of calcium from the surface and gelation of alginate. The hydrogel coated scaffolds were then incubated in APS (9 mg/mL) and TEMED (6 mg/mL) in DI water solution for 5 min to crosslink the PEGDMA macromonomers within the hydrogel layer and form an interpenetrating network. The scaffold/hydrogel composite was washed with DI water to remove the residual initiator or unreacted macromonomers.
Coating Non-Porous Rods with a Bioresorbable Hydrogel
The procedure for coating non-porous PCL-TCP rods with a bioresorbable hydrogel is schematically shown in
For freeze drying, the HyTECs were submerged in liquid nitrogen for 30 min and then lyophilized and stored at 4° C. For rehydration, the freeze-dried HyTECs were incubated in DI water for 15 min before further analysis. For E-beam sterilization, HyTECs were exposed to E-beam irradiation at a standard single dose of 25 kGy, following norm ISO 11137-2:2006. E-beam sterilization of BMP2-laden hydrogel-loaded PCL-TCP rods was performed with a single dose (25 kGy) or two (12.5 kGy) doses to see the effect of splitting E-beam dose on the activity of loaded protein.
The density of grafted APMA on PCL-TCP scaffolds was quantified via measuring the concentration of unreacted APMA in the solution after the reaction using ninhydrin assay. Briefly, the APMA solution after the reaction with PCL-TCP scaffold was diluted 10 times in MES buffer. 40 μL of the ninhydrin reagent was added to 200 μL of the diluted APMA solution. After mixing, the solution was heated to 90° C. in a capped tube for 8 min and the absorbance was read at 570 nm using a SpectraMax M2 plate reader (Molecular Devices LLC). The concentration of unreacted APMA in the solution was calculated using a calibration curve made for the absorbance of solutions with known concentrations of APMA. To evaluate the surface hydrophilicity of PCL-TCP constructs, a 4 μL water droplet was deposited on the disc scaffolds, the contact angle was measured using a goniometer Ramé-Hart 290 (Ramé-Hart instrument co., USA) and analyzed using image processing.
The apparent Young's modulus and the stress at yield of the scaffolds were tested using an Instron 5944 uniaxial testing system (Instron Corporation, Norwood, MA) with a 2 kN load-cell, a preload of 1N and a displacement rate of 1% strain/s. The initial slope of the stress vs strain curve was taken as the Young's modulus. The stress at yield was defined as the stress at which a line starting from 1% strain offset with a slope equal to the Young's modulus intersected with the stress vs strain curve.
To measure the release of calcium ions from the surface of the scaffolds, the CaCl2 treated scaffolds (or CaSO4 treated non-porous rods) were incubated in 1 mL DI water at room temperature for 1 hr. The concentration of Ca2+ ions in the release medium was measured using calcium colorimetric assay (MAK022, Sigma-Aldrich, USA) on a SpectraMax M2 plate reader (Molecular Devices LLC) at 575 nm.
To measure the hydrogel layer thickness, the scaffolds were imaged before and after hydrogel loading using a Dino-Lite digital microscope camera. The images were then analyzed using ImageJ to quantify the average hydrogel thickness. The average thickness of the gel in the quadrilateral pores was defined as half of the difference between the size of the empty spot (black color in images) before and after the hydrogel loading. The fraction of filled pores was defined as the ratio of the number of those quadrilateral pores that were completely filled with hydrogel to all quadrilateral pores.
The hydrogel loading (%) was calculated from the scaffold weight before hydrogel loading (Wb) and after hydrogel loading (Wa), using the following equation;
For scanning electron microscopy (SEM) imaging, the HyTEC samples were immersed in liquid nitrogen and freeze-dried. The freeze-dried samples were dipped in liquid nitrogen and cut using a surgical blade. The hydrogel samples were then coated with gold using a SPI sputter (SPI Supplier Division of Structure Prob, Inc., West Chester, PA) for 180 seconds and imaged using a Field Emission Scanning Electron Microscope (Zeiss Sigma, White Plains, NY) at an accelerating voltage of 5 keV.
A customized 3D printed PCL-TCP device was designed and used to evaluate the adhesion of the hydrogels to PCL-TCP scaffolds (see
For measurement of release kinetics from porous HyTECs, BMP2 protein was added to the PEGDMA (10-30% wt/vol), alginate (1.5% wt/vol) and APS (1 mg/mL) precursor solution prior to hydrogel loading on 80% porous PCL-TCP scaffolds. The average hydrogel loading (relative to the scaffold weight) changed from 151% to 169% and 144% with increasing the PEGDMA concentration from 10% to 20% and 30% (see
For measurement of BMP2 release kinetics from non-porous HyTECs, the rod-shaped HyTECs with 2 μg encapsulated BMP2 were freeze-dried and incubated in 1 mL PBS at 37° C. for 28 days. At each time point, the amount of BMP2 in the release medium was measured using ELISA and the release medium was replaced with fresh PBS.
Human Mesenchymal Stem Cells (hMSCs) and Multi-potent mouse C3H10T1/2 fibroblasts (ATCC, USA) were cultured in DMEM medium (Life Technologies, USA) supplemented with 10% fetal bovine serum (FBS, Life Technologies, USA) and 1% Penicillin and Streptomycin (hereafter referred to as culture medium) at 37° C. in a 5% CO2 humidified incubator. After reaching 70% confluency, hMSCs or C3H10s were enzymatically lifted with trypsin-EDTA and used for in-vitro studies. All cells were passaged <6 times prior to the in-vitro studies.
For in-vitro cell studies, the 80% porous PCL-TCP scaffolds after APMA surface modification and before CaCl2) treatment, were sterilized in 70% ethanol solution for 20 min.
The CaCl2) solution and hydrogel precursor solution were sterilized by filtration using 0.22 μm Millex syringe filters.
The scaffolds with or without hydrogel loading were incubated in 1 mL culture medium at 37° C. The culture medium with no scaffold exposure incubated at 37° C. was used as the control group. At days 1 (for viability and proliferation tests) and 4 (for proliferation test) the incubated cell culture medium (hereafter referred to as conditioned medium) was used for viability and proliferation tests.
For cell viability test, concurrent with incubation of scaffold or scaffold/hydrogel in culture medium, hMSC and C3H10 cells were seeded on 96 well plates at 5000 cells/well and incubated at 37° C. and 5% CO2 for 24 hr. The cultured medium was then replaced with 100 μL of the conditioned culture medium and the cells were incubated for another 24 hr. To measure the cell viability, 10 μL of the CCK-8 solution (CCK-8 kit, Sigma-Aldrich) was added to each well and after 3 hours of incubation, the absorbance was read at 450 nm on a plate reader. The viability of cells in the experimental groups (scaffold, scaffold+gel) was divided by that of cells in the control group (no scaffold) to calculate the normalized viability.
For cell proliferation test, concurrent with incubation of scaffold or scaffold/hydrogel in culture medium, hMSC and C3H10 cells were seeded on 24 well plates at 10000 cells/well and incubated at 37° C. and 5% CO2 for 24 hr. The cultured medium was then replaced with 600 μL the conditioned culture medium and the cells were incubated for 7 days. 3 days after addition of conditioned medium (4 days after cell seeding), the medium was replaced with fresh conditioned medium. At days 0, 3 and 7, cells were washed with PBS, enzymatically lifted using 250 μL of 0.25% trypsin-EDTA solution (Life Technologies, USA) and counted using a Z2 particle counter (Beckman Coulter, USA).
The HyTECs without loaded BMP2 (scaffold+gel) or with 1.5 μg loaded BMP2 (scaffold+gel/BMP2), were incubated in 1 mL culture medium at 37° C. The conditioned medium was used for cell differentiation test every 3 days and replaced with fresh culture medium. The culture medium with no scaffold and no BMP2 (ctrl), culture medium supplemented with 1.5 μg/mL BMP2 (BMP2 in medium (3 d)) and culture medium supplemented with 214 ng/mL BMP2 for 21 days (BMP2 in medium (21 d)) incubated at 37° C. were used as control groups. Concurrent with incubation of experimental and control groups in culture medium, hMSC and C3H10 cells were seeded on 24 well plates at 10000 cells/well and incubated at 37° C. and 5% CO2 for 24 hr. The cultured medium was then replaced with 600 μL of the conditioned culture medium and the cells were incubated for 21 days with changing the medium to fresh conditioned medium every 3 days. For BMP2 in medium (3 d) group, the medium was changed to culture medium (BMP2 free) after 3 days. At each time point (day 0, 7, 14 and 21), cells were washed with PBS and lysed with 1% Triton X-100 in PBS using a cell scraper followed by shaking for 20 min at room temperature. The lysate was centrifuged at 2000×g for 15 min at 4° C. and the supernatant was collected. The ALP activity in the supernatant was measured using QuantiChrom ALP Assay Kit (BioAssay Systems, Hayward, CA, USA) according to the manufacturer's Instructions, on the plate reader at 405 nm. The double-stranded DNA content of the lysate was measured using PicoGreen assay kit (Quant-it, Thermo Fisher Scientific). The ALP activity was divided by the DNA content to calculate the normalized ALP activity.
PCL-TCP filaments with 0.9 mm diameter were synthesized and coated with GelMA as described in the previous section. Then, the GelMA coated rods were dipped into a CaSO4 suspension in DI water (100 mg/mL) at 60° C. and sonicated for 30 seconds. The rods were then transferred into wells of a 24-well plate, dried under vacuum, and sterilized under UV for 60 minutes. hMSCs were suspended in hydrogel precursor solution containing GelMA (10%), Alginate (1.25%), PEGDMA (2%), and photoinitiator (0.3%) in calcium-free culture medium at 2 million cells/mL density. The sterile rods were then dipped in the cellular precursor solution for 2 minutes at 37° C. The cell-laden hydrogel-loaded rods were removed from the solution, left in sterile dry wells of a 96-well plate for 5 minutes, and then irradiated with visible light for 15 minutes. The cell-laden hydrogel coated rods were then transferred into wells of a 24-well plate and incubated in culture medium at 37° C. and 5% CO2.
For live/dead cell imaging, cell-laden HyTECs were stained with Calcein AM (2 μM) and Ethidium homodimer-1 (4 μM) according to manufacturer's instructions and imaged using a Zeiss AxioObserver Z1 fluorescent microscope. The live/dead images were divided into smaller squares and the number of live and dead cells were counted manually to calculate the cell viability. To quantify the DNA content of the HyTECs, at each time point, the samples were transferred into new wells and incubated in 500 μL of DMEM medium supplemented with collagenase (1 mg/mL) for 1 hour at 37° C. Then, 250 μL of triton solution (3%) in PBS was added to each well and the attached cells were scrapped from the surface using a CytoOne cell scraper (USA Scientific Inc, Ocala, FL). Then the cell suspension was transferred to a microcentrifuge tube and sonicated. The cell lysate was then centrifuged at 2000×g at 4° C. for 15 min and the supernatant was collected. The content of double-stranded DNA in the supernatant was measured using Quant-iT PicoGreen DNA assay according to manufacturer's instructions.
All experiments were done in triplicate. Statistically significant differences between groups were tested using a two-way ANOVA with replication, followed by a two-tailed Students t-test. A p-value smaller than 0.05 (p<0.05) was considered statistically significant.
The density of grafted APMA on PCL-TCP constructs versus APMA concentration in the reaction solution is shown in
The effect of CaCl2) concentration in the incubation solution on the hydrogel coating of 80% porous PCL-TCP scaffolds is shown in
3D printed porous PCL-TCP scaffolds with 50%, 60% and 70% porosity, before and after hydrogel coating are shown in
The effect of freeze drying on the mechanical properties of PCL-TCP scaffolds and characteristics of the hydrogel layer on the scaffolds are shown in
The structure of the 3D printed PCL-TCP device which was used to measure the adhesion of hydrogels to scaffolds is shown in
The normalized viability and proliferation of hMSC and C3H10 cells cultured in DMEM medium which was preconditioned with PCL-TCP/hydrogel or pristine PCL-TCP scaffolds are shown in
The ALP activity of hMSC and C3H10 cells cultured in DMEM medium (ctrl), DMEM medium preconditioned with scaffold/hydrogel without BMP2 (scaffold+gel), DMEM medium preconditioned with scaffold+BMP2 loaded hydrogel (scaffold+gel/BMP2), DMEM medium supplemented with 1.5 μg/mL BMP2 for 3 days (no scaffold/BMP2 (3 d)), and DMEM medium supplemented with 214 ng/mL BMP2 for 21 days (no scaffold/BMP2 (21 d)) are shown in
For making non-porous HyTEC constructs, NaOH treatment followed by freezing/thawing, GelMA conjugation to the surface, and CaSO4 deposition were used to increase hydrophilicity, improve hydrogel adhesion and stimulate surface-initiated crosslinking, respectively. An SEM image of the surface of the hydrogel loaded on non-porous PCL-TCP rods is shown in
The complications associated with autografts, allografts and DBM for treatment of large bone defects highlights the importance of developing synthetic bone grafts. A number of studies have shown PCL-TCP scaffolds are biocompatible, bioresorbable, mechanically stable and osteo-conductive. Further, owing to the low melting point and processability of PCL, PCL-TCP scaffolds with well-controlled porosity can be manufactured using Fused Deposition Modeling (FDM)-based 3D printing. However, lack of bone growth stimulating proteins limits the application of 3D printed PCL-TCP scaffolds for treatment of large bone defects. In the present invention, the inventors developed a postprocessing method for manufacturing interconnected porous, protein-laden thick hydrogel layer-coated 3D printed PCL-TCP scaffolds. Following 3D printing, the surface of scaffolds was treated in three consecutive steps to increase hydrophilicity, improve hydrogel adhesion and stimulate surface-initiated crosslinking. NaOH treatment imparts hydrophilicity to the surface of polyesters due to the scission of ester bonds to carboxyl and hydroxyl groups. Reactive double bonds were then incorporated onto the surface by grafting APMA to carboxyl groups using carbodiimide chemistry. To control the hydrogel layer thickness, the APMA modified scaffolds were treated with CaCl2 solution to stimulate a surface-initiated crosslinking. When the CaCl2 treated scaffolds were dipped into hydrogel precursor solution, the deposited CaCl2) diffused from the surface to the solution, crosslinked alginate at the proximity of the surface, and made a hydrogel layer on the surface. The PEGDMA macromonomers within the physically crosslinked hydrogel were covalently crosslinked in the next step to form a stiff interpenetrating network. The hydrogel network bound to the scaffold surface through reaction of double bonds of PEGDMA macromonomers and double bonds of APMA grafted to the scaffold surface. The covalent binding between reactive functional groups of PEGDMA and those of scaffold surface during the crosslinking reaction, increased the hydrogel adhesion to the surface of PCL-TCP scaffolds. Adhesion of polymer networks to rigid surfaces increases when functional groups on the surface link to the polymer network. For instance, the adhesion between a PEGDA/alginate IPN hydrogel and glass, ceramics, titanium or aluminum significantly increased with modifying the surface with reactive 3-(trimethoxysilyl) propyl methacrylate (TMSPMA) and covalent anchoring the hydrogel to the surface.
The hydrogel layer thickness was directly correlated with the total amount of released calcium ion from the surface and could be tailored with altering processing parameters including NaOH treatment time and CaCl2) concentration in the treatment solution. An increase in the total amount of released calcium ion from the scaffolds with raising NaOH treatment time was due to an improved surface hydrophilicity and roughness, hence higher absorption of CaCl2 solution on the PCL-TCP surface. Likewise, an increase in the total amount of released calcium ions from the scaffolds with raising CaCl2) concentration in the treatment solution was due to a larger calcium deposition on the surface.
Despite an elevated hydrogel loading, the hydrogel thickness did not dramatically change with increasing the scaffold porosity. Therefore, the pore size of porous HyTECs could be tuned by changing the scaffold porosity. The size of interconnected pores in bone tissue engineering scaffolds should be at least 100 μm for cell infiltration, bone ingrowth, and vascularization. However, cell migration and bone ingrowth is optimal when the pore size was larger than 100 μm. For example, the adhesion and proliferation of osteoblasts on collagen-glycosaminoglycan (CG) scaffolds with pore size greater than 300 μm were higher than those scaffolds with pore size smaller than 200 μm [24]. When porous poly(ether ester) block-copolymer scaffolds were implanted into the dorsal skinfold chamber of balb/c mice, vessel ingrowth was faster for the scaffolds with large pores (250-300 μm) compared to those with medium pores (75-212 μm) or small pores (20-75 μm). In this invention, the scaffolds pore size after hydrogel loading ranged from 140 μm to 300 μm, 480 μm, 1100 μm when the porosity of pristine scaffolds increased from 50% to 60%, 70% and 80%. Although the pore size of all hydrogel loaded scaffolds was larger than 100 μm, those scaffolds with minimum 60% porosity and 300 μm or larger pore size after hydrogel loading would be optimal for future in-vivo experiments, based on the aforementioned published reports.
Results presented herein showed that BMP2 was released from alginate/PEG-based hydrogel loaded porous PCL-TCP scaffolds over 35 days. It has been demonstrated that a BMP2 loaded alginate/PEG based hydrogel with a sustained release of BMP2 in-vitro stimulated ectopic bone nodule formation in mice. Therefore, incorporation of BMP2-laden alginate/PEG-based hydrogel imparts osteo-inductivity to osteoconductive 3D printed porous PCL-TCP scaffolds and potentially enhances and accelerates bone formation. A lower amount of released protein at higher PEGDMA concentrations (
Freeze-drying facilitates storage/transportation and increases the shelf-life of bioactive products. Results presented herein showed that freeze-drying did not adversely influence the mechanical properties of the scaffolds and characteristics of the hydrogel layer but reduced the release of bioactive BMP2. PCL-TCP scaffolds are heat sensitive and E-beam irradiation is considered a reliable method for terminal sterilization of heat sensitive materials. It was also shown that E-beam sterilization did not adversely affect mechanical properties and degradation kinetics of PCL-TCP scaffolds. The results of the present invention revealed that E-beam sterilization reduced the release of bioactive BMP2 from freeze-dried BMP2-laden porous HyTECs.
The results presented herein further showed that osteogenic differentiation of hMSCs and C3H10 cells was higher when the cells were exposed to BMP2 released from porous HyTECs compared to 3-day exposure of cells to BMP2 dissolved in the medium at the same dose as BMP2 loading in hydrogel. The higher osteogenic differentiation of cells exposed to BMP2 releasing scaffolds compared to 3-day exposure of cells to BMP2 was due to the time-dependent osteo-inductivity of BMP2. Osteo-inductivity of BMP2 protein is dose- and time-dependent. For instance, the ALP activity of hMSCs exposed to slow BMP2 releasing electrospun PCL/PEG mats was significantly higher than that of hMSCs exposed to fast BMP2 releasing mats.
The method of this invention could be used for loading a wide range of hydrogels on porous or non-porous polyester-based constructs. Therefore, in addition to the porous scaffolds, the inventors investigated the efficacy of the method for making non-porous HyTECs with a bioresorbable hydrogel for sustained release of BMP2 protein. To improve the integration of hydrogel with the non-porous rod, freezing/thawing was used after NaOH treatment to increase the surface roughness. The results showed that freezing/thawing increased calcium deposition on the surface of the rods (not shown). GelMA and PEGDMA were used as macromonomer and crosslinker, respectively. HepMA was used for prolonging the release of BMP2, due to a high affinity of heparin to BMP2. It has been shown that addition of HepMA to an alginate-based hydrogel extended the release kinetics of BMP2 and improved subcutaneous bone formation in mice. In addition, photo-initiation was used for covalent crosslinking of the hydrogel on non-porous rods instead of chemical-initiation that was used for crosslinking of hydrogels in porous scaffolds. Chemically initiated crosslinking was used in porous scaffolds because the penetration of light into the central parts of the scaffold might be limited. The release of BMP2 (as a model protein) from non-porous HyTECs showed that the method can be used for sustained delivery of proteins along with non-porous implants. Total amount of released enzymatically active BMP2 from hydrogel loaded rods after freeze-drying and single E-beam sterilization was only 28%. That might be attributed to denaturation of BMP2 or crosslinking of a GelMA based network with BMP2 under highly intense radiation. However, the amount of released active BMP2 almost doubled with splitting a highly intense E-beam dose (25 kGy) to two doses (12.5 kGy) with lower intensity. The inventors also showed that the method described herein could be used for loading live cell-laden hydrogels on scaffolds. The biodegradable hydrogel layer was thick enough to accommodate cells and the cells were viable and proliferating. The described method could be particularly useful for cell loading on implants.
The present invention claims the benefit, or priority, to U.S. Provisional Applications 63/289,431 filed Dec. 14, 2021, 63/304,216 filed Jan. 28, 2022, 63/289,447 filed Dec. 14, 2021, and 63/304,207 filed Jan. 28, 2022 all of which are incorporated herein by reference for all that they teach.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2022/052407 | 12/9/2022 | WO |
| Number | Date | Country | |
|---|---|---|---|
| 63289431 | Dec 2021 | US | |
| 63304216 | Jan 2022 | US | |
| 63289447 | Dec 2021 | US | |
| 63304207 | Jan 2022 | US |
