WO2024091971A1 - 3d printed gene-activated scaffolds for bone regeneration - Google Patents

Abstract

Described herein are compositions and methods for bone regeneration and the treatment of bone diseases, bone defects, and injuries to bone. The compositions and methods include treating bone regenerative implants with lyoprotectants to enhance their surface and mechanical properties and then gene- activate the resulting scaffolds to enhance bone regeneration.

Description

3D PRINTED GENE-ACTIVATED SCAFFOLDS FOR BONE REGENERATION

Cross-reference to Related Application

This application claims the priority of U.S. provisional application Serial No. 63/418,793, filed October 24, 2022, the disclosure of which is incorporated herein by reference in their entirety as if fully set forth herein.

Background

Tissue engineering is an interdisciplinary field aimed at restoring or improving tissue function in damaged or dysfunctional tissues by regenerating newly functional tissue, usually with the help of a biomaterial scaffold. Tissue engineering can be used to describe a variety of tissue regeneration techniques, however, the term is most commonly used to describe the creation of cell-material constructs and the culturing of these constructs in vitro or their implantation in vivo. Researchers in the field are applying tissue engineering techniques to the treatment of tissue-related injuries and diseases in a variety of tissues and organs, such as bone, skin, heart, and liver. Given the vast functional, structural, and mechanical differences between these tissues, each target tissue requires a different approach to properly replicate the native tissue and restore its original function.

A variety of biomaterials have been investigated as scaffolds to guide the regeneration of tissue towards a desired morphological or functional end, including collagen, calcium phosphate, synthetic polymers, and combinations thereof. To date, the simplest tissue engineering endeavors (i.e. those aimed at the simplest tissues) have been the most successful, with tissue engineered products aimed at regenerating dermal skin, bone, and cartilage having already gained regulatory approval from the Food and Drug Administration in the United States. However, tissue engineering has had difficulty with replicating the intricate structural features key to kidney and liver function and has not come close to producing the organ-replacement therapeutics that have been the long-term goal of the field since its inception. Summary

Compositions and methods are described herein that are useful for bone regeneration and the treatment of bone diseases, bone defects, and injuries to bone. The compositions and methods include use of bone regenerative implants treated to enhance their surface and mechanical properties and then gene-activate the resulting scaffolds to enhance bone regeneration.

It has been demonstrated that orthopedic implants integrate with bone better when the implant's surface is highly textured. It has also been shown that implants intended to regenerate bone perform better when the implant has a highly textured surface. The creation of large-volume implants for bone regeneration is an active area of research that has not yet reached a consensus on the best implant composition or design for bone regenerati on. Despite a lack of consensus, studies investigating 3D printed bone regenerative implants have shown that creating a pore network throughout the entire volume of a bone regenerative implant improves bone regeneration outcomes. Studies that have investigated using highly textured, 3D printed, porous implants for bone regeneration applications have found promising results. However, the ability of water to penetrate into the center of the implant can be hindered by the highly textured surface of the implant. Materials with microtextured and nanotextured surfaces have been found to exhibit water repelling behavior that is caused by the interaction between water and the highly textured surface (termed the "lotus effect").

Compositions and methods described herein provide methods for allowing water-based solutions to enter into the internal pores of any porous 3 D implant but is especially applicable to implants that have microtextured or nanotextured surfaces. The method enables the loading of a variety of drugs, delivery systems, or other components into a bone regenerative implant. In embodiments, a product of the method was used to load a gene-delivery system into implants thereby allowing the implant to induce production of a therapeutic protein at the site of implantation.

Methods are also described herein that involve making porous implants that allows water to penetrate the entire implant. For example, methods includes treating a bone implant, such as implants made from calcium phosphate cement (CPC implants) with a lyoprotectant, e.g., a 2% to 10 %, e.g., 4%, w/v% solution of sucralose, sucrose or trehalose in pure water, exposing the solution containing the implant to vacuum, e.g., to remove the bubbles trapped inside the implant's pore network and replaces them with lyoprotectant solution (vacuum defoaming), removing the implant from the solution and freezing the implant, and then lyophilizing (freeze dry) the implant. The result is an implant with its pores filled with a substance resembling cotton candy, termed "lyophilized cake". When a water-based solution is added to an implant treated as described above, the cake conducts the solution through the pores and allows the internal volume of the implant to be filled with the solution. The resulting filled implant can then be used as-is or lyophilized again to preserve any materials contained in the solution for long-term storage. The method thus improves the functionality of porous 3D implants, especially implants that have micro/nanotextured surfaces and orthopedic implants aimed at regenerating bone.

Treating implants with the method described above can allow for loading with any water-based solution, like solutions containing gene-delivery reagents, small molecule drugs, or protein therapeutics. The method allows for conduction of water through the pores of the scaffold using a hydrophilic material suspended in the void space (volume) between surfaces in contrast to surface coating that covers the surface with hydrophilic materials to improve the wetting of that surface by masking the hydrophobic surface underneath.

Description of the Figures

Figs. 1A-1C are illustrations of different methods of 3D printing. Fig. 1A depicts extrusion printing. Fig. IB depicts powder bed fusion printing. Fig. 1C depicts stereolithography (SLA) printing.

Fig- 2 is schematic illustrating the mechanisms of transfection from scaffolds activated with pDNA, viruses, interfering RNAs, and RNA transcripts.

Fig. 3 is a schematic illustrating BMP signalling that leads to osteogenic differentiation.

Figs. 4A and 4B are images of calcium phosphate cement (CPC) meshes showing the precision of printing in the macroscale (Fig. 4A) and within the mesh’s internal pore network (Fig. 4B).

Fig. 4C is a 3 -dimensional illustration of a CPC implant.

Figs. 5A-5H are SEM images of printed CPC constructs at varying magnifications. Fig. 5A shows an image of a 3D CPC mesh with 0.2 mm strand diameter and 50% infill at 30X magnification, wherein the CPC mesh is treated with water alone. Figs. 5B, 5C, and 5D are the same as Fig. 5A, but treated with vapor alone, vapor + water, or vapor + SBF, respectively. Fig. 5E-5H are the same treatments as Figs. 4A-4D at 10,000x magnification. At 30X magnification, “Water Alone” meshes display cracks in the strands, while other groups do not. At lOkX magnification, the “Vapor Alone” surfaces appear to have less surface texture than the other groups.

Figs. 6A-6C are images of CPC 3D print types used for CPC characterization experiments. Fig. 6A is an image of CPC 3D discs. Fig. 6B is an image of CPC 3D cylindrical meshes. Fig. 6C is an image of CPC 3D rods.

Fig. 7 is a graph showing the specific surface area of printed CPC rods (0.2 mm diameter, 4 mm length) as calculated by the Brunauer-Emmet-Teller (BET) method using nitrogen gas as the adsorbate. Bars represent mean ± SD, n=3-4. Significant differences between samples were assessed using ordinary oneway ANOVA and Tukey’s multiple comparisons tests (*p<0.05, **p<0.01, ****p<0.0001).

Figs. 8A and 8B are graphs showing compressive strength (Fig. 8A) and Young’s modulus (Fig. 8B) of 3DP CPC cylindrical meshes printed with 0.2 mm strand diameter, 50% infill, 7.5 mm total diameter, and 3 mm height. Bars represent mean ± SD, n=6 for all groups. Significant differences between samples were assessed using Brown-Forsythe and Welch tests with Dunnett’s T3 multiple comparisons tests (*p<0.05, **p<0.01).

Figs. 9A and 9B are graphs showing calcium content (Fig. 9A) and pH (Fig. 9B) of medium incubated with CPC discs for 24 hr. Bars represent mean ± SD, n=3 for all groups. Significant differences were assessed using ordinary oneway ANOVA with Tukey’s multiple comparisons tests (*p<0.05, **p<0.01, ***p<0.001, ****p < 0.0001).

Fig. 10 is a graph showing the viability of HEK 293T cells seeded onto CPC discs for 48-96 hours. CPC discs were either untreated (CPC) or had been incubated in complete DMEM for 24 hours prior to cell seeding (pre-incubated CPC). Bars represent mean ± SD, n=3 for all groups. Significant differences between samples were assessed using unpaired t-tests (**p<0.01, ****p<0.0001).

Fig. 11 is a series of SEM images of CPC prints at increasing levels of magnification to illustrate the relative size of CPC surface texture as compared to HEK 293T cells. Fig. 12 is a graph showing Alkaline phosphatase activity from cell lysate of BMSCs cultured on CPC discs for 7 days (normalized to total DNA). Bars represent mean ± SD, n=5 for all groups. Significant differences between groups were assessed using ordinary one-way ANOVA with Tukey’s multiple comparisons tests (*p<0.05, **p<0.01, ***p<0.001, **** pO.OOOl).

Figs. 13A-13C show HEK 293T cells treated with lyophilized polyplexes in solutions with final sucrose concentrations of 1-10% (w/v) and reconstituted these solutions in ultrapure water. Fig. 13A shows fluorescence microscopy images of HEK 293T cells 48 hours after transfection with polyplex solutions containing 2.5 pg of EGFP pDNA and containing sucrose in a range of concentrations (1-10%, w/v) that were lyophilized in microfuge tubes and reconstituted. Scale bar represents 1000 pm. Fig. 13B is a graph depicting 48 hours transfection efficiencies of HEK293T cells transfected with either plain polyplex solution (0% Sucrose) or polyplex solution containing 1% sucrose (1% Sucrose) that had each been lyophilized on a titanium disc. Fig. 13C is a graph depicting log transformed mean fluorescence of HEK 293T cells transfected with the polyplex solutions described above. All cells transfected via disc exposure were treated with polyplexes containing 5.5 pg of pDNA. Values are expressed as mean ± SD (n=3). Significant differences between samples were assessed using one-way ANOVA and Tukey’s multiple comparisons tests (****p<0.0001). Laird et al.

Figs. 14A-14F are graphs showing transfection efficiencies and log transformed mean fluorescence 48 hours after transfection of cells with titanium discs coated with 1% (w/v) sucrose polyplexes containing indicated amounts of EGFP pDNA. Figs. 14A and 14D show HEK293T cells. Figs. 14B and 14E show primary gingival keratinocytes. Figs. 14C and 14F show primary gingival fibroblasts. Values are expressed as mean ± SD (n = 3). Significant differences were assessed using one-way ANOVA with Tukey’s multiple comparisons tests (*p<0.05, **p<0.01, ***p<0.001, ****p<0.0001). Laird et al.

Figs. 15A and 15B are graphs showing average particle diameter and poly dispersity index (Fig. 15A) and zeta potential (Fig. 15B) of polyplexes before and after lyophilization with varying sucrose concentrations. Values are expressed as mean + SD or mean ± SD (n=4). Significant differences between samples’ particle diameter and zeta potential were assessed by unpaired t-tests (ns: no significance, *p<0.05, **p<0.01, ***p<0.001). Laird et al.

Fig. 16 is a graph showing transfection efficiency of HEK 293T cells 48 hours after transfection with polyplexes that were freshly prepared or lyophilized in 2% sucralose and stored for 24 hours, 1 year, or 2 years. Values are expressed as mean ± SD, n = 4-6. Significant differences between groups were assessed using ordinary one-way ANOVA with Tukey’s tests (*p<0.05, ****p<0.0001). Malkawi el al.

Figs. 17A-17C show transfection efficiency of cells treated with the supernatant of the CPC/polyplex mixture and cells treated with the resuspended pellet of CPC powder. Fig. 17A shows composite micrographs (fluorescence and brightfield) of HEK 293T cells treated with either water (Untreated), polyplex solution (Polyplexes), centrifuged polyplex solution (Polyplex Supernatant), CPC powder mixed with polyplex solution (CPC/Poly Mixture), or the supernatant from the centrifugation of the previous mixture (CPC/Poly Supernatant). Bright dots indicate fluorescence from EGFP. Fig. 17B is a graph showing transfection efficiency for HEK 293T cells treated as described above. Fig. 17C is a graph of log transformed mean fluorescence intensity for HEK 293T cells treated as described above. Cells were incubated for 72 hours after transfection. Bars represent mean + SD, n=3 for all groups. Significant differences were assessed using ordinary one-way ANOVA with Tukey’s multiple comparisons tests (****p<0.0001).

Figs. 18A-18C show that CPC discs gene-activated by the lyophilized coating method with sucralose (Lyo. Loaded CPC) or the adsorption method (Ads. Loaded CPC) successfully transfected cells when seeded with HEK 293T cells. Fig. 18A shows fluorescence micrographs of HEK 293T cells seeded onto CPC discs in well plates and exposed to polyplexes. CPC discs were either: untreated (CPC Negative Control), incubated in polyplex solution for 1 hour (Ads. Loaded CPC), or had polyplex/lyoprotectant solution pipetted onto them before lyophilization (Lyo. Loaded CPC). Positive control groups had polyplex solution added directly after seeding, negative control groups had water added directly after seeding. Bright spots indicate fluorescence from cytosolic EGFP. Transfection efficiencies (Fig. 18B) and mean fluorescence intensities (Fig. 18C) for HEK 293T cells seeded onto CPC discs or tissue culture plastic (TCP) and exposed to polyplexes as described above. Cells were incubated for 72 hours after seeding. Bars represent mean + SD, n=3 for all groups. Significant differences were assessed using ordinary one-way ANOVA with Tukey’s multiple comparisons tests (ns = no significance, *p<0.05, **p<0.01, ***p<0.001, ****p < 0.0001).

Figs. 19A and 19B are graphs showing the transfection efficiencies of gene-activated CPC discs hardened with four different hardening methods. Fig. 19A shows transfection efficiencies of HEK 293T cells seeded onto CPC discs that were hardened via the indicated methods and gene-activated via the Lyo. Coat method. Fig. 19B shows log transformed mean fluorescence intensity of HEK 293T cells treated as described above. Cells were incubated for 72 hours after seeding. Bars represent mean ± SD, n=9-10. Significant differences were assessed using either one-way ANOVA with Tukey’s multiple comparisons tests (A) or the Kruskal -Wallis test with Dunn’s multiple comparisons tests (B) (*p<0.05, **p<0.01, ***p<0.001).

Figs. 20A-20C show increasing transfected cells of the same mesh over 96 hours. Fig. 20A is a graph showing transfection efficiency 96 hours after HEK 293T cells were seeded onto CPC scaffolds loaded with polyplexes (Gene- Activated Scaffold) or treated with either freshly made polyplexes (Fresh polyplexes) or polyplexes made at the time of scaffold gene-activation (Lyophilized Polyplexes). Fig. 20B is a graph of log transformed mean fluorescence intensity of HEK 293T cells treated as described above. Fig. 20C shows fluorescence micrographs of transfected HEK 293T cells after incubation for 48 hours, 72 hours, and 96 hours. Bright spots are fluorescence from EGFP. Bars represent mean ± SD, n=3-4. Significant differences between groups were assessed using either ordinary one-way ANOVA with Tukey’s multiple comparisons tests (A) or Brown -Forsythe and Welch tests with Dunnett’s T3 multiple comparisons tests (B) (ns: no significance, **p<0.01, ***p<0.001, ****p<0.0001).

Figs. 21A-21E are images of untreated and treated calcium phosphate cement (CPC) implants that are highly textured at the nanoscale. Fig. 21A is a side view of a water droplet placed on top of the untreated CPC implant and unable to absorb. Fig. 21B is a side view of a CPC implant treated with lyoprotectant after having absorbed water. Fig. 21C is a top view of the untreated CPC implant with the unabsorbed water droplet. Fig. 21D is a side view of a CPC implant treated with lyoprotectant after having absorbed water. Fig. 21E is a schematic of a Wenzel and Cassie-Baxter staters of liquid-surface interfaces.

Fig. 22 is a schematic illustration of a method of treatment of CPC implants with a lyoprotectant.

Fig. 23 is a series of images depicting water being continuously pipetted onto treated porous CPC implant, which readily absorbs all water pipetted onto it.

Fig. 24 is an image of histological sections of bone showing changes in contact osteogenesis from use of an untreated CPC implant (upper panels) and a CPC implant treated with a lyoprotectant-containing solution (lower panels).

Detailed Description

Compositions and methods are described herein that are useful for bone regeneration to treat of bone diseases, bone defects, and injuries to bone. The compositions described herein can be administered to treat subjects, such as animals or humans, in need of such treatment, or who can develop a need for such treatment. For example, the compositions can repair, replace, and/or regenerate bone that has been damaged from trauma (e.g. automobile accidents, explosive incidents), congenital bone malformations, debridement of infected bone tissue, bone cancer tumor resections, or osteoporosis-related fractures. The bone can be any bone in the subject’s body including facial bones, skeletal long, short, flat, or irregular bones. The subject’s bone(s) can be treated upon an initial trauma to the bone, during healing, or after healing.

Bone tissue has the ability to heal itself naturally, however, that natural healing can be insufficient if the defect is too large (in humans, typically when the defect is >2.5 cm in length). Such defects that are not expected to heal over the remainder of the patient’s lifetime are termed “critical-sized defects”, and are typically treated with an autograft, allograft, or a non-biological implant. While these approaches can be successful for both congenital defects and injuries, they can also have drawbacks such as limited autograft volume, risk of disease transmission, risk of immune reaction, and long-term implant failure. Bone tissue engineering aims to match the therapeutic efficacy of autografts while not being limited by graft availability constraints and, when compared to allografts, also avoid the risks of rejection and disease transmission. The goal is to create a therapeutic that can: regenerate bone within the defect; grow and change within pediatric patients; and does not require a secondary surgery site for donor bone collection (as is the case for autografts).

Researchers in the field of bone tissue engineering often use a biomaterial scaffold to induce and/or guide bone regeneration. Through their work they found that endowing a scaffold with certain properties yields superior bone regeneration outcomes in vivo. Specifically, a rough exterior surface, ^43-46 an interconnected pore network,^47-49 and pore sizes of 2OO-4OOpm^50-53 throughout the volume of the scaffold have each been shown to improve bone regeneration in scaffolds regardless of the material used.

3D Printing for Bone Tissue Engineering

Three-dimensional printing (3DP) has become popular within tissue engineering because the technique allows precise control over material composition and morphology. 3DP is a manufacturing technique characterized by using a printing device to systematically add material to a printing surface to assemble a digitally designed object. The most common 3DP methods are material extrusion, powder bed fusion, and stereolithography, which all use a standardized method of assembly in which an object is created by stacking a series of thin layers on a printing surface. In material extrusion printing (also known as fused deposition modeling, or FDM), fluidic material is extruded in a defined pattern onto the previous layer to create a stack of layers that form the printed object (Fig. 1A). In powder bed fusion printing (commonly known as selective laser sintering, or SLS), an even layer of particulates is fused together (via laser sintering or addition of a binding liquid) in a defined pattern on top of the prior layer to create a stack of layers that form the printed object (Fig. IB). In stereolithographic (SLA) printing, an inverted print surface is immersed in a vat of curable liquid (often a photopolymerizable polymer solution) that is selectively cured in a layer-by-layer fashion to form the printed object (Fig. 1C). Methods for printing a specific object depend on factors such as the size of the object, the type of material desired, and the desired resolution.

3DP can be used to dictate the structure of the internal pore network within bone implant scaffolds by adjusting the pore diameter, pore density, and pore shape in the software used to design their scaffolds. An additional benefit to using 3DP is that a scaffold can be designed to perfectly replicate most 3D shapes, and 3D medical imaging techniques can be combined with 3DP to create scaffolds that perfectly match a patient’s own morphology. This means that a 3D printed bone regenerative therapeutic might be used in craniofacial reconstruction and plastic surgery to regenerate bone that exactly replicates the patient’s original bone structure or even amends a pre-existing congenital defect.

In vitro characterization of 3D printed calcium phosphate cement (CPC) scaffolds

Using 3DP to create scaffolds for tissue engineering allows the production of scaffolds with complex and patient-specific 3D macroscale morphologies, but the macroscale morphology of a scaffold is just one of its many important characteristics. For successful tissue engineering, a scaffold comprises a material that can support and/or stimulate formation of the desired of tissue. For bone regeneration, calcium phosphate cement (CPC) is an attractive printing material due to its ability to stimulate differentiation toward bone cells and its mechanical strength. In general, CPCs are cementitious materials comprised of fine calcium phosphate particulates mixed with a carrier fluid. The exact composition of the calcium phosphate particulates can vary, with some formulations including particulates of specific calcium phosphate phases, such as hydroxyapatite or P- tricalcium phosphate. The carrier fluid also varies, with some groups using water or aqueous gels, while others use an anhydrous mixture of various oily liquids. The CPCs can be hardened through a dissolution and reprecipitation process during prolonged exposure to water at room temperature. This sets CPC-based scaffolds apart from other calcium-phosphate based scaffolds that use sintering for hardening because sintering results in a much smoother surface texture on the scaffold. This is important because researchers investigating bone cell differentiation on the surface of implants made of titanium, polymers, and calcium phosphate have all shown that having greater surface roughness results in greater bone cell differentiation. The Ginebra group at the Polytechnic University of Catalonia has shown that the morphology of the nanotexture of CPC materials also impacts bone healing independent of the surface’s average roughness and calcium content of the surrounding cell culture medium. As a result of these findings, the composition herein comprise CPC to print scaffolds.

CPC can be manufactured by Innotere GmBH that uses an oily carrier liquid instead of water. The CPC formulation from Innotere has been well- characterized by the Gelinsky group at the Technical University of Dresden. Treatment of the CPC formulation by immersion in water and incubation in a highly humid environment (for 3 days in both cases) both serve to transform the CPC into calcium deficient hydroxyapatite. The hardening protocols for CPC scaffolds are described below.

Gene delivery

Researchers have shown that incorporating a gene delivery system into scaffolds can increase their functionality by inducing cells local to the scaffold and/or cells that invade the scaffold to produce therapeutic proteins that enhance healing outcomes. “Gene delivery” as used herein is a process that includes (1) the delivery of nucleic acids (2) encoding a gene of interest (3) to a recipient organism’s cells without modification of the cells’ genomes (4) for the purpose of inducing transient expression of that gene in those cells.

The goal of gene delivery (as defined above) is to induce transient expression of genes of interest for a therapeutic goal. As illustrated in Fig. 2, the mechanisms of transfection from scaffolds activated with nucleic acids such as plasmid DNA (pDNA), interfering RNAs, RNA transcripts, and viruses have multiple routes. RNA transcripts have the simplest transfection mechanism (Route A), where they can be translated into osteogenic protein immediately after entering the cytosol. Interfering RNAs follow a different transfection mechanism (Route B), where after entering the cytosol they can inhibit the anti-osteogenic activity of targeted mRNAs and miRNAs. Viruses commonly used in gene-based therapy and pDNA follow similar routes within the cell (Route C), where after endosomal escape they must reach the nucleus. After entering the nucleus, the encoded genes can be transcribed into mRNA that is then translated into osteogenic protein.

An example of a successful gene delivery therapeutic are the mRNA-based CO VID- 19 vaccines, which have been immensely useful for bringing the CO VID- 19 pandemic under control. In the case of the COVID-19 vaccines, gene delivery was used to induce expression of a specific viral protein to train recipients’ immune systems to target the COVID-19 virus. However, gene delivery can also be used to induce a patient’s cells to produce therapeutic proteins that would otherwise be administered as part of a medical device or a drug treatment regimen. In this way, a patient’s body becomes the manufacturing site of its own medicine and as such this process could be an alternative to producing the proteins through an expensive and time-consuming biologic drug manufacturing process.

An example of the effectiveness of including a gene delivery system for bone regeneration can be demonstrated by the induction of gene expression to produce bone morphogenetic protein 2 (BMP -2) (signaling pathway depicted in Fig- 3), among other osteogenic proteins. BMP-2 is already used in the clinic to enhance bone regeneration, but it is used in its protein form. The protein itself is very potent and does indeed enhance bone healing but administering BMP-2 protein by loading it onto scaffolds can be problematic. To maintain a therapeutic concentration of BMP-2 protein at the injury site, large amounts of the protein is loaded onto the scaffold to combat the dilution, degradation, and aggregation that all proteins experience in vivo. The large doses required mean that therapeutics containing BMP-2 are very expensive and they can lead to off-target effects, including ectopic bone formation in local non-bone tissue. By using a gene delivery approach, off-target effects are reduced or prevented and costs are reduced.

Within gene delivery there are two main classes of delivery vectors: viral gene delivery vectors and non-viral gene delivery vectors. As described herein, non-viral vectors are used because of the challenges associated with viral vectors. Specifically, viral vectors are known to induce potentially hazardous immune responses. For example, Egermann et al. described the hazards of using viral vectors in bone regeneration applications when they showed that using an adenoviral vector to stimulate BMP-2 production in a large animal bone regeneration model produced an immune response to the viral vector that induced systemic suppression of bone regeneration.

An illustration of BMP signaling that leads to osteogenic differentiation is shown in Fig. 3. BMPs induce Smad-dependent and non-Smad-dependent signaling. In the Smad-dependent pathway, Smad (1, 5, or 8) is phosphorylated, complexes with Smad 4, then translocates to the nucleus where co-factors are recruited (including Runx2) for the regulation of osteogenic gene expression. In the non-Smad-dependent pathway, TAK1 recruits TAB1 to stimulate activation of MKK3/6 which in turn activates p38a/p. p38a/p phosphorylates Runx2, Dlx5, and Osx, all of which regulate osteogenic gene expression. BMP: Bone morphogenetic protein; BMPR: BMP receptor. Wu el a!.. adapted in Laird et al. A bone regenerative therapeutic that combines 3DP and gene delivery in a single device to combine their benefits is provided. To do this, 3D printed scaffolds were made from calcium phosphate cement with defined pore networks, their surfaces were then optimized for bone regeneration, and the scaffolds were gene-activated by loading them with PEI-pDNA polyplexes.

The invention will be described by the following non-limiting examples.

Example 1: Methods

Printing of CPC scaffolds:

Non-aqueous CPC was purchased from Innotere GmbH (Germany) and loaded into 3 mL printing cartridges (Nordson EFD, Westlake, OH, USA) affixed with a 27G conical needle tip (Fisnar, Germantown, WI, USA). The cartridges were loaded into a BioX 3D printer (Cellink, Sweden), and printed using a 3 mL pneumatic printhead (Cellink). Printing was performed at room temperature with printing pressures of 100-700 kPa and printing speeds of 3-10 mm/s. All scaffold designs were created using 3D Builder (Microsoft, Redmond, WA, USA) and sliced using Slic3r (version 1.3.0, open source).

Hardening and texturing of CPC scaffolds:

After printing, scaffolds were either immersed in ultrapure water at room temperature (“water alone method”, or “WA”) or placed in a humidified incubator at 37°C (“vapor alone method”, or “VA”) for 3 days. After the incubation, scaffolds were washed 3 times with acetone (20 minutes of immersion in acetone per wash), then left to dry at room temperature. Some of the scaffolds hardened with the VA method were immersed in either ultrapure water (“vapor + water”, or “VW) or SBF (“vapor + SBF”, or “VS”) for 3 days at room temperature. SBF was prepared by dissolving salts (all from Sigma Aldrich) in ultrapure water (ThermoFisher Scientific) to produce a solution containing 40.286 mM NaCl, 1.143 mM KCl, 0.143 mM MgSCU, 0.286 mM MgCl, 1.2mM NaHCO₃, 2.5 mM CaCh, and 1 mM KH2PO4. CPC scaffolds were incubated in SBF at a volume/weight ratio of 21.52 mL of SBF per 1 g of CPC. After the secondary incubations, the acetone washes were repeated then the scaffolds were left to dry at room temperature.

Scanning electron microscopy:

Cylindrical mesh scaffolds (7.5 mm in diameter and 1 mm in height, 50% rectilinear infill pattern, 0.2 mm strand diameter) were printed and hardened as described above. Samples were sputter coated with a gold-palladium alloy and imaged with field emission scanning electron microscopy (Hitachi S-4800, Hitachi, Japan).

Compression testing:

Cylindrical mesh scaffolds (7.5 mm in diameter and 3 mm in height, 50% rectilinear infill pattern, 0.2 mm strand diameter) were printed and hardened as described above. Scaffolds were compressed with an MTS Insight Material Testing System (MTS Systems, Eden Prairie, MN, USA) loaded with a IkN load cell at a rate of 1 mm/min until failure. The resulting stress vs. strain plots were used to calculate compressive strength and Young’s modulus.

Specific surface area analysis:

Rods 0.2 mm in diameter and 2 mm long were printed and hardened as described above. Rods were divided into sample groups of equal weight, then analyzed with a Beckman-Coulter surface analyzer (Model SA-3100, Miami, FL). N2 adsorption isotherms at 77K were collected over a relative pressure range of 0.00 to 0.20 in 0.02 increments. The Brunauer, Emmett, and Teller (BET) theory was applied to five points in the relative pressure range of 0.00 to 0.20. All data analyses are built into the SA-3100 analyzer.

Calcium content and pH after incubation:

Discs (7.5 mm in diameter, 1 mm in height, 100% infill, no distinction between strands) were printed and hardened as described above. Each disc was incubated in 0.5 mL of Dulbecco's Modified Eagle Medium (DMEM) containing 1% sodium pyruvate, 1% HEPES buffer, 1% Glutamax (all from ThermoFisher Scientific, Waltham, MA, USA), 0.05 mg/ml gentamycin sulfate (IBI Scientific, Dubuque, IA, USA), and 10% fetal bovine serum (FBS, Atlanta Biologicals, Flowery Branch, GA, USA) (termed “complete DMEM” hereafter) in a humidified incubator for 24 hr. Medium was removed and aliquots were removed for pH measurement (S20 SevenEasy, Mettler Toledo, Columbus, OH, USA). Calcium content was assessed with a commercial calcium colorimetric assay kit (MAK022, Millipore Sigma, Burlington, MA, USA) according to the manufacturers protocol.

Cell viability on CPC:

Discs (7.5 mm in diameter, 1 mm in height, 100% infill, no distinction between strands) were printed and hardened with the VA method, as described above. A portion of the discs were incubated in 0.5 mL of complete DMEM in a humidified incubator for 24 hours (“pre-incubated CPC”) while the rest were stored (“CPC”). All discs were then immersed in 0.5 mL of complete DMEM in a 48-well plate, then seeded with 37,500 HEK 293T cells obtained from the American Type Culture Collection (ATCC, Manassas, VA, USA). Discs and cells were incubated for the requisite time, then MTS cell proliferation assay reagent (CellTiter 96® AQueous One Solution Cell Proliferation Assay, Promega, Madison, WI, USA) was used to determine viability of HEK 293T cells. Specifically, 100 uL of MTS assay reagent was added to each well (20 uL per 100 uL of medium), the mixture was incubated according to the manufacturer’s protocol, then 120 uL aliquots were removed from each well and their absorbance at 490 nm was measured with a microplate reader (SpectraMax Plus 384, Molecular Devices, San Jose, CA, USA). Absorbance values were normalized with the absorbance of untreated cells at each timepoint to yield relative viability values.

Alkaline phosphatase activity assay:

Discs (7.5 mm in diameter, 1 mm in height, 100% infill, no distinction between strands) were printed and hardened as described above. Each disc was incubated in 0.5 mL of complete DMEM in a humidified incubator for 24 hours prior to having the medium replaced and being seeded with 12,500 BMSCs obtained from the American Type Culture Collection (ATCC). Discs and cells were incubated for 7 days, after which an alkaline phosphatase activity assay (ab83369, Abeam, United Kingdom) was used according to the manufacturer’s protocol to assess the activity in the cell lysate. Total DNA from the cell lysate was measured via a PicoGreen assay (ThermoFisher Scientific) according to the manufacturer’ s protocol and was used to normalize the enzyme activity data. Statistical analysis:

All statistical analysis was performed with Prism (version 9, GraphPad, San Diego, CA, USA). If statistical comparisons were to be made between more than two groups within a dataset, the datasets were assessed with QQ plots, residual normality tests, and homoscedasticity tests to check whether the ANOVA assumptions of equal variance and normal distribution were violated. If a dataset failed the assumption of normal distribution, the Kruskal -Wallis test with Dunn’s multiple comparisons testing was used for statistical analysis. If a dataset passed the assumption of normality but failed the assumption of equal variance, then the Brown-Forsythe and Welch test with Dunnett’s T3 multiple comparisons testing was used for statistical analysis. If a dataset passed both the assumption of normal distribution and the assumption of equal variance, then ordinary one-way ANOVA with Tukey’s multiple comparisons testing was used for statistical analysis. The statistical tests used for each dataset are indicated in the figure legend.

Example 2: Scaffold hardening method

CPC scaffolds were printed at a resolution of 200 pm using a pneumatic 3D printer (Fig. 4A-4B). An illustration of the CPC scaffold is shown in Fig 4C. Eventually a standard approach was developed: set the printhead speed to lOmm/s and then manually adjust the printing pressure during printing to maintain a continuous flow of material. The printing pressure was highly variable (with a range of 100-700kPa) because CPC is viscous and moves through the printing cartridge in an inconsistent manner. Specifically, the pneumatic pressure applied to the CPC often formed a uniquely shaped channel within the material that created a variable surface area for the pneumatic pressure to act against. As a result, the pressure required to maintain a continuous flow of material varied with the shape of the channel created each time a new cartridge of CPC was loaded for printing.

Once the printing of CPC had been worked out, two methods of hardening the CPC scaffolds were investigated that had been reported in the literature already: (1) immersing the scaffolds in water for 3 days (termed “water alone”, or “WA”) and (2) incubating the scaffolds in a highly humid environment for 3 days (termed “vapor alone”, or “VA”). Both methods had already been confirmed to adequately harden the CPC in earlier work, but this earlier work noted that the WA method caused cracks to form in the printed strands that led to weaker mechanical properties, while the VA method did not produce cracks (Fig. 5A and 5B). However, this prior work had not rigorously investigated the effect of these methods on the nanotexture of the CPC material. Scanning electron microscopy (SEM) was used to image the CPC scaffolds hardened by both the WA and VA methods and stark differences were observed in the surface texture. Specifically, the WA method produced a highly textured surface filled with nanoscale plate- like structures (Fig. 5E), while the VA method produced a smoother surface free of the large plate-like structures (Fig. 5F).

A hardening procedure was then developed that would prevent the crack formation seen with the WA method, but still give the final scaffold a nanotextured surface. We investigated adding an additional incubation step after hardening the scaffolds with the VA method. Specifically, the VA-treated scaffolds were incubated in ultrapure water (termed “vapor + water”, or “VW”) or simulated body fluid (SBF) (termed “vapor + SBF”, or “VS”) for an additional 3 days. We selected SBF as an incubation medium because of its known ability to stimulate mineralization of materials incubated in such solutions. Additional incubations did not introduce cracks to the scaffolds (Fig. 5C and 5D) but produced a surface that appeared more textured than the surfaces produced with the VA method (Fig. 5G and 5H). Rather than label one surface as more textured than the other based on images, the material surfaces were quantitatively compare by using the Brunauer-Emmett-Teller (BET) method for specific surface area (SSA) analysis. SSA analysis consists of measuring the total surface area of a sample, then normalizing it to the weight of the sample. The SSA for a sample of particulates can be used as a proxy for the particulates’ surface roughness so long as the particulates are similar in volume. To ensure the samples were uniform, we printed a large number of 2 mm long rods (shown in Fig. 6C), hardened them using our four hardening methods, then measured the SSA of the resulting rods with the BET method.

The WA group had the highest SSA (34.072 m²/g) and the VA group had the lowest SSA (10.259 m²/g), while the VW and VS groups were in between (14.8107 m²/g and 13.597 m²/g, respectively). The VW and VS groups were not significantly different from each other, but they were both significantly different from the WA and VA groups (Fig. 7). We were unsure whether to trust the SSA value for the WA group because the WA method creates cracks in the rods, which opens up more surface area for nitrogen to adsorb to during the SSA analysis (Fig. 5A). Since the nitrogen adsorbed to the surface exposed by the cracks is not distinguishable from the nitrogen adsorbed to the exterior surface, the SSA value for the WA group may be artificially inflated by nitrogen adsorbing to the surface exposed by the cracks. Thus, the results may not show a significantly greater roughness for the WA group relative to the other groups. Since the VW and VS groups both appeared more textured in images and had greater SSA values than the VA group, the additional incubations demonstrated increased the roughness of the surfaces.

Comparison of mechanical strength:

Earlier work has shown that the cracks produced with the WA method make the scaffolds mechanically weaker than scaffolds prepared with the VA method. As a result of this observation, the mechanical properties of scaffolds produced with our new VW and VS methods were assessed. Mechanical compression testing were performed on 3 mm by 7.5 mm cylindrical meshes (Fig. 6B) to determine their compressive strength (maximum load before failure) and their Young’s modulus (resistance to being compressed). Both compressive strength and Young’s modulus the mean values were found for each group followed the same pattern: WA<VA<VW<VS (Figs. 8A and 8B). The VS scaffolds were significantly different from both the WA and VA scaffolds in both metrics, though the Young’s modulus for the VW scaffolds was significantly different from that of the WA scaffolds. From this we concluded that the VS scaffolds have stronger compressive mechanical properties than the other groups. Comparison of effect on cell culture medium:

As mentioned above, the CPC is known to transform into calcium deficient hydroxyapatite during hardening. While this phase of calcium phosphate is similar to hydroxyapatite (the primary mineral component of natural bone), the qualifier “calcium deficient” means that the material will absorb calcium from its surroundings until it becomes “stoichiometric” hydroxyapatite. Specifically, when calcium deficient hydroxyapatite is immersed in a solution containing calcium (such as cell culture medium), calcium from the surroundings will be absorbed, phosphate will be released, and the medium will become more basic as a result of the released phosphate. This phenomenon has been linked to a decline in the viability of cells incubated with calcium deficient hydroxyapatite materials. Based on these findings, calcium content and pH of medium was measured after 24 hours of incubation with 1 mm by 7.5 mm discs (Fig. 6A) to assess the ion flux caused by our scaffolds. We found that our data was similar to the literature, with all CPC scaffolds reducing the calcium content to less than 50% of the content of DMEM and increasing medium pH relative to DMEM (Figs. 9A and 9B). The magnitude of calcium reduction and pH increase varied between the CPC hardening methods, with the VA scaffolds having both the greatest decline in calcium content and the greatest increase in medium pH. We believe this occurred because the VA scaffolds were not immersed in any liquid medium at any time during preparation, and thus the ions within the scaffold were unable to dissolve and reprecipitate into more thermodynamically stable calcium phosphate crystals. When the scaffolds were immersed in cell culture medium, these less thermodynamically stable crystals may have been more soluble, leading to greater ion flux. However, the increase in pH observed after incubation of CPC materials should enhance bone formation, since it has been reported that slightly alkaline environments enhance bone cell activity while acidic conditions impair bone cell activity.

We also assessed whether incubating our CPC constructs in cell culture medium could reduce the decline in viability of seeded cells. CPC discs were prepared with the VA hardening method, incubated some of them in complete DMEM (termed “pre-incubated CPC”) for 24 hr, then seeded them with cells, and assessed viability over 96 hours. We found that the pre-incubated CPC discs had improved viability compared to the untreated CPC at all timepoints (Fig. 10), but that the viability was still less than cells not exposed to CPC. This apparent toxicity is inferred to be due to the ion flux from calcium absorption by CPC based on published research demonstrating that calcium ion flux is a major factor in toxicity in cells seeded onto CPC materials.

Comparison of effect on osteogenic differentiation of bone marrow-derived mesenchymal stem cells:

Differences in bone marrow-derived mesenchymal stem/stromal cells (BMSC) differentiation induced by the morphology of the nanotexture on CPC surfaces has been reported. Specifically, “needle-like” nanocrystals outperform “plate-like” nanocrystals in terms of BMSC osteogenic differentiation and bone regeneration in vivo. Unfortunately, no mechanistic explanation has been offered, though it may be mediated by the interaction of cells with the nanoscale texture on the surface of CPC materials (shown in Fig. 11).

The surfaces produced with our hardening methods were compared to determine whether the differing surface morphologies observed could produce a difference in BMSC differentiation. BMSCs were seeded onto the discs and incubated for 7 days, after which the cells were lysed and their alkaline phosphatase (ALP) activity and total DNA content were assessed. ALP activity is an early indicator of osteogenic differentiation, and is a common assay used to assess osteogenic differentiation in vitro. ALP activities from our different hardening methods were found to follow the same pattern as the mechanical testing results: WA<VA<VW<VS (Fig. 12). Interestingly, while the VW and VS groups were not significantly different from each other and both showed significant differences from the WA and VA groups, the VS group showed stronger levels of significant difference from the WA and VA groups. From this we concluded that the differing hardening methods do result in differences in differentiation, with the VS method seeming to produce the most differentiation.

In sum, textural differences in our CPC scaffolds were observed after using different hardening methods, confirmed those observed differences with SSA analysis, then differences in mechanical properties and osteogenic differentiation were demonstrated to be a result of the different hardening methods. We also showed that our material behaves like other CPC constructs reported in the literature by absorbing calcium from surrounding medium and modifying the pH through release of phosphate. Incubating the CPC constructs in cell culture medium to allow ion exchange prior to cell exposure was also shown to reduce the toxic effect on cells from the ion flux. Based on these characterization experiments, the VS hardening method was demonstrated to be an effective candidate for producing CPC scaffolds intended for bone regeneration.

Example 3: Formulating lyophilizable gene delivery systems for use on calcium phosphate cement scaffolds:

The hardening methods developed for our CPC scaffolds improved the mechanical properties and the osteogenicity of our scaffolds, but we wanted to stimulate bone regeneration even more. In prior work, the Salem lab has investigated loading non-viral gene delivery systems onto scaffolds made of collagen to enhance their bone regenerative capabilities.

Complexation of plasmid DNA (“pDNA”) with cationic polymers to form polyplexes is a non-viral gene therapy approach. For example, pDNA can be complexed with polyethylenimine (PEI) to produce “PEI-pDNA polyplexes”. PELpDNA polyplexes are formed by mixing cationic PEI with anionic pDNA in solution at specific charge ratios, after which the two components spontaneously complex into polyplexes. These polyplexes have been shown to be endocytosed by cells and decomplex, allowing pDNA that has undergone nuclear translocation to induce expression of the gene it encodes. The exact mechanisms by which this decomplexation and nuclear translocation occur are yet to be fully understood, but the fact that the polyplexes are endocytosed and that the pDNA can reach the nucleus in some way is well established (schematic shown in Fig. 2, Route C). PEI-pDNA polyplexes were loaded onto scaffolds, a process termed “geneactivation”, but this time the gene-activating was occurring on a scaffold made of calcium phosphate rather than collagen. To do this, two gene-activation methods were tested, one of which was a gene-activation method that originally developed for use on titanium dental implant surfaces. The intent for gene-activating these dental implant surfaces was to induce the expression of proteins that would enhance the soft tissue seal around the implant and prevent bacterial downgrowth. While the original goal of that project was to gene-activate a dental implant surface, the gene-activation method we developed for it can be used to gene- activate just about any implant surface due to its simplicity. The gene-activation method consists of preparing a solution of PEI-pDNA polyplexes, mixing in a lyoprotectant, pipetting the mixture onto the implant surface, then lyophilizing the coated implant.

Our initial work investigated using sucrose as the lyoprotectant for coating since sucrose is a very common lyoprotectant. However, using sucrose as part of a treatment to prevent bacterial downgrowth on a dental implant can be counterproductive, since sucrose is a major contributor to the growth of bacteria in the oral environment and is known to contribute to cavity formation because of the bacterial growth it enables. As a result of this, the amount of sucrose present on the dental implant surface was reduced as much as possible by reducing the concentration of sucrose in the polyplex solution. However, this approach still resulted in sucrose being added to the oral environment, so we decided to remove sucrose entirely. We reasoned that replacing sucrose with a molecule that has a similar structure to that of sucrose could work well as a replacement, since the mechanism of sucrose’s lyoprotection is based on its structure, but such a molecule would have to be non-metabolizable to circumvent any concerns of promoting bacterial growth. We found that sucralose, a non-metabolizable chlorinated sucrose derivative marketed as “Splenda”, was an excellent candidate, and selected this molecule as our lyoprotectant for the gene-activation of our CPC scaffolds. We also compared this gene-activation method to an adsorption method that we developed based on results showing adsorption of polyplexes to CPC surfaces.

Results and Discussion:

Sucrose-based coatings of titanium discs:

As described above, the first lyophilizable gene delivery system we investigated consisted of a mixture of sucrose and polyplex solutions that were then lyophilized. Since sucrose is cariogenic, we tried to reduce the amount of sucrose present in the solution as much as possible by reducing the final concentration of sucrose. We lyophilized polyplexes in solutions with final sucrose concentrations of 1-10% (w/v), reconstituted these solutions in ultrapure water, then treated HEK 293T cells with the polyplexes. We found that even a final concentration of 1% sucrose (29.2 mM) was able to preserve transfection activity after reconstituting the lyophilized product with pure water (Fig. 13A). We then confirmed that the sucrose was required for transfection to occur after lyophilization (Figs. 13B and 13C) and demonstrated that an immortalized cell line (HEK 293T) and two types of oral primary cells (human gingival fibroblasts and human gingival keratinocytes) could be transfected in a dose-dependent manner after a titanium surface coated with varying doses of polyplexes lyophilized in 1% sucrose was suspended above the cells (Fig. 14). The polyplexes were released from the lyophilized cake upon exposure to aqueous medium (in this case cell culture medium), after which they could transfect local cells. The polyplex size, size distribution (poly dispersity index), and surface charge (zeta potential) was characterized before and after lyophilization in varying concentrations of sucrose. We found that the samples containing no sucrose and 1% sucrose had elevated average particle diameter after lyophilization and much greater poly dispersity indices (Fig. 15A). However, assessment of the polyplexes’ surface charge before and after lyophilization showed that only the 0% (w/v) sucrose concentration had a significant shift in surface charge after lyophilization (Fig. 15B)

Transition to sucralose as lyoprotectant: As mentioned above, replacing sucrose with a non-cariogenic alternative would better suit the oral environment since stimulating bacterial growth through the introduction of any amount of sucrose would be counterproductive. Sucralose was investigated as an alternative compound because of its structural resemblance to the potent lyoprotectant sucrose and because of its non-metabolizable nature. The structures of sucrose and sucralose is shown below with chlorine substitutions in sucralose indicated by red arrows.

Sucrose Sucralose

Sucralose was found to function as a lyoprotectant for PEI-pDNA polyplexes in an initial study (data not shown). A long-term polyplex preservation study with sucralose was then performed. We found that polyplexes lyophilized in 2% (w/v) sucralose were able to transfect cells up to two years after preparation while retaining -66% of the total transfection activity of polyplexes that were prepared the day of treatment (Fig. 16). Interestingly, after lyophilization the polyplexes exhibited only a small decline in transfection efficiency after two years of storage, with 24 hour old polyplexes yielding a transfection efficiency of 44.6% while 2 year old polyplexes yielded a transfection efficiency of 40.1%. It is important to note that shortly after we began investigation of sucralose as a lyoprotectant, the Grohganz group at the University of Copenhagen assessed the chemical properties of a variety of excipients to determine whether they could perform as lyoprotectants and identified sucralose as a possible candidate, among many others. They then tested 64 of the identified compounds, including sucralose, as lyoprotectants for protein solutions and found that sucralose outperformed all tested lyoprotectants in the 12 protein solutions they tested. We decided to use sucralose as the lyoprotectant for future gene-activation studies because we believed lyophilized sucralose would be more stable than lyophilized sucrose due to sucralose being non-hygroscopic while sucrose is hygroscopic.

Gene-activation of CPC constructs: The adsorption of polyplexes to a CPC surface was then investigated. Since CPC has an ionic surface due to it being comprised of calcium and phosphate ions, the charged PEI-pDNA polyplexes might adsorb to the surface via ionic interactions and still be endocytosed by cells. To test the ability of polyplexes to adsorb to CPC, we prepared some polyplexes and homogenized some CPC prints to create CPC powder. The CPC powder was mixed with a portion of the polyplexes, after which both the CPC/polyplex mixture and an aliquot of the polyplex solution were centrifuged for 30 s at 10,000 g. HEK 293T cells were then treated with the resulting supernatants (“Polyplex Supernatant”, “CPC/Poly Supernatant”). The pelleted CPC powder was then resuspended and cells were treated with the resulting mixture (“CPC/Poly Mixture”) or an uncentrifuged aliquot of polyplexes (“Polyplexes”). Assessment of transfection via fluorescence microscopy and flow cytometry both showed that centrifugation of polyplex solutions without CPC powder does not result in a reduced transfection efficiency, implying that the polyplexes are unaffected by the centrifugation process (Fig. 17). However, cells treated with the supernatant of the CPC/polyplex mixture showed no transfection while cells treated with the resuspended pellet of CPC powder displayed significant albeit marginal transfection (Fig. 17). This suggests that the polyplexes exposed to CPC powder were removed from the solution during centrifugation via adsorption to CPC particulates, and that some of those polyplexes were able to transfect cells treated with the particulates.

Since polyplex adsorption to CPC yielded detectable transfection, we chose to compare this adsorption method to the lyophilized coatings that we developed for the titanium dental implant surfaces. CPC discs were prepared, hardened with the VA method, then gene-activated via either the lyophilized coating method with sucralose (Lyo. Loaded CPC) or the adsorption method (Ads. Loaded CPC). When seeded with HEK 293T cells, both gene-activated CPC discs successfully transfected cells (Fig. 18), but the Lyo. Loaded CPC showed more transfection from cells that landed on the tissue culture plastic of the well plate during seeding (Fig. 18A). Transfection of cells local to the CPC, but not adhered directly to the CPC, was expected based on the transfection experiments with titanium discs, wherein the coated titanium discs were suspended above cells seeded onto tissue culture plastic in well plates. While the Ads. Loaded CPC discs had a lower mean transfection efficiency and log(mean fluorescence) than the Lyo. Loaded CPC discs, the values were not significantly different (Fig. 18B and 18C). We had some concerns about the shelf-stability of CPC gene-activated with the adsorption method: Leaving the CPC in polyplex solution would enable continuous dissolution/precipitation of the calcium and phosphate ions that would alter the nanotexture in unpredictable ways and drying the CPC to prevent the nanotexture alteration could inactivate the polyplexes. Additionally, it was thought to be beneficial for the polyplexes not to be bound to the surface of the scaffold and instead released into the surrounding medium upon immersion in medium, as they were with the lyophilized coatings for titanium discs. As a result of these concerns and a desire for immediate release, the lyophilized coating method for all future experiments.

The different hardening methods were then tested to produce different transfection efficiencies after gene-activation with the lyophilized coating method, so we gene-activated CPC discs hardened with our 4 hardening methods and tested their ability to transfect HEK 293T cells (Fig. 19). We found that there was a difference in the transfection efficiency between the hardening methods, with the VW and VS methods producing the highest mean transfection efficiency and mean fluorescence intensity, albeit with high variability (Fig. 19A and B).

After comparing the transfection efficiencies of gene-activated CPC discs with different hardening methods, we decided to confirm that a 3D mesh made of CPC could be successfully gene-activated since thus far we had only shown transfection from the 2D surface of CPC discs. A cylindrical 3D mesh was prepared. It was hardened using the VS method, and gene-activated via an adapted version of the lyophilization coating method. The scaffolds were seeded with HEK 293T cells by pipetting cell suspensions onto dry meshes, observed the cells with fluorescence microscopy over 4 days (96 hr), then assessed transfection efficiency with flow cytometry. We found that transfection occurred almost exclusively in cells in direct contact with the mesh (Fig. 20C), implying that nearly no polyplexes were released into the cell culture medium after the HEK 293 T cell suspension was pipetted onto the meshes. This outcome is different from the immediate release of polyplexes that occurred in the lyophilized coatings of titanium surfaces and resulted in strong transfection after just 48 hours (Fig. 13 and Fig. 14), and is likely explained by the observation from our prior experiments showing that polyplexes adsorb to the CPC surface. However, the transfection efficiency for cells seeded onto the mesh was much lower than any transfection efficiency observed from any of the CPC discs or titanium discs, regardless of hardening method or gene-activation method (0.95% transfection for the 3D meshes vs. 2.3%-13.4% transfection for the CPC discs and 47.3% transfection for similar doses on titanium discs).

This low transfection may be due to the method used to seed cells onto the meshes. Specifically, in this experiment a cell suspension of HEK 293T cells was vigorously pipetted directly on top of a mesh in a well plate, then left to incubate without any additional agitation or inversion of the mesh. We expected this method to be sufficient for our experimental purposes because we expected the polyplexes to be trapped in the lyophilized cake and then be released by dissolution of the cake after exposure to aqueous fluids, as was the case for our lyophilized coatings of titanium dental implant surfaces. A significant transfection in cells adhered to tissue culture plastic was also observed when performing transfection studies with CPC discs (Fig. 18A). We did not think that the polyplexes had been incubated with the mesh long enough prior to freezing (~5 minutes) to facilitate major adsorption of our polyplexes to the CPC surface. However, since we did not observe any transfection in cells adhered to tissue culture plastic, it is probable that during this short exposure period nearly all the polyplexes were adsorbed to the surfaces of the mesh, including the lateral sides and underside of each strand in the mesh. This positioning of polyplex adsorption is important to note because in early work we noticed that cells in suspension seeded onto CPC 3D meshes seemed to roll off the rounded strands of the meshes and onto the tissue culture plastic the mesh was resting on. In fact, this observation informed our choice to use CPC discs (with flat surfaces for cells to land on) for subsequent cell-based experiments. Since the cell suspension was vigorously pipetted onto the CPC 3D mesh and since there was no subsequent agitation or inversion during the initial steps of incubation, it is likely that only a small proportion of the seeded cells adhered to the meshes while the rest landed on tissue culture plastic. If we assume that (nearly) all polyplexes were adsorbed to the mesh’s surface, then transfection should occur (almost) exclusively by cell interaction with the mesh’s surface. As a result, we should expect the level of transfection observed to be proportional to the surface area of the mesh that is covered by cells, which should increase during incubation as cells continue replicating and migrating across the surface of the mesh. This type of gradual increase in transfection from gene-activated scaffolds has already been observed in studies that seeded cells onto gene-activated scaffolds containing hydroxyapatite. We may have also observed this gradual increase in transfection over time, since observation of the same mesh over 96 hours showed increasing numbers of transfected cells (Fig. 20C). However, these images do not depict the transfection efficiency of the cells on the meshes since we cannot visualize untransfected cells, and so we must declare this explanation as merely a possibility and not a probability due to a lack of quantitative data. In future in vitro studies, we should use cell seeding protocols that allow for the robust and even seeding of cells across the entire surface of our scaffolds, such as the protocol described in Wash et al. (2021).

In sum, PEI-pDNA polyplexes were shown to be formulated as lyophilizable coatings for implant surfaces by addition of sucrose and subsequent lyophilization. Sucralose was then shown to function as a lyoprotectant of PEI- pDNA polyplexes and lyophilization in sucralose can preserve polyplex function for at least two years. We then applied this gene-activation method to CPC scaffolds and determined that CPC materials can also be gene-activated via adsorption of polyplexes to the surface. After deciding to move forward with the lyophilized coating method of gene-activation and not the surface adsorption method, VW and VS hardening methods were determined to produce more transfection than the WA and VA methods. We then tested the lyophilized coating method of gene-activation with a 3D printed mesh made of CPC, and found that while it can transfect cells, it seems as if the lyophilized coating method was not distinct from the adsorption method. This is because only the cells adhered to the CPC got transfected, implying adsorption of all polyplexes loaded onto the CPC meshes with the lyophilized coating method. Based on the sum of these results, we determined that we could gene-activate CPC scaffolds and moved forward with testing bone regeneration in vivo using our lyophilized gene-activation method.

Methods:

Purification of plasmid DNA:

Plasmid DNA was purified from DH5a Escherichia coli that had been previously transformed. Purification was performed using a GenElute HP Endotoxin-Free Plasmid Maxiprep Kit (Sigma-Aldrich, St. Louis, MO, USA) according to the manufacturer's protocol.

Polyplex solution preparation:

Polyplexes were prepared as described previously.¹¹² Briefly, two 500 pl solutions containing either 130 pg of 25-kDa branched PEI (Sigma-Aldrich) or 100 pg of pDNA were prepared in DNAse/RNAse-free water (ThermoFisher Scientific). The PEI solution was added to the pDNA solution, vortexed for 30s, and incubated for 30 minutes to allow for complexation between the pDNA and PEI. The resulting 1ml polyplex solution had a nitrogen (N) to phosphate (P) ratio (N/P ratio) of 10, which was shown to yield maximum transfection efficiency with minimal cytotoxicity. Varying volumes of 40% sucrose or 20% sucralose in DNAse/RNAse-free water (ThermoFisher Scientific) were added to the polyplex solution to yield the desired sucrose or sucralose concentrations.

Polyplex characterization:

Polyplexes were prepared as described above, then frozen in a -80°C freezer and lyophilized (FreeZone 4.5 -105, Labconco, Kansas City, MO, USA). Their zeta potential and hydrodynamic size (pre-lyophilization and postlyophilization) were then measured via dynamic light scattering and electrophoretic light scattering with a Zetasizer Nano-ZS (Malvern Instruments, UK) according to the manufacturer's protocol.

Cell culture:

HEK 293 T cells, primary gingival fibroblasts, and primary gingival keratinocytes were obtained from the American Type Culture Collection (ATCC, Manassas, VA, USA). Cells were cultured in complete Dulbecco's Modified Eagle Medium (DMEM) containing 1% sodium pyruvate, 1% HEPES buffer, 1% Glutamax (all from ThermoFisher Scientific), 0.05-mg/ml gentamycin sulfate (IBI Scientific, Dubuque, IA, USA), and 10% fetal bovine serum (FBS, Atlanta Biologicals, Flowery Branch, GA, USA) in a humidified incubator at standard culture conditions (37°C and 5% CO2, Sanyo Scientific, Japan). Cells were passaged with 0.25% trypsin-EDTA (ThermoFisher Scientific).

Preparation and coating of titanium discs with polyplexes:

Titanium discs (12 mm in diameter, 3 mm in height) made of commercially pure titanium were prepared as described previously.^{117 133} Briefly, discs were sanded with a variable speed grinder-polisher (Ecomet 3, Buehler, Lake Bluff, IL, USA) using grinding papers (CarbiMet, Buehler) ascending to grit number 600. The titanium discs were then sandblasted (EWL Type 5423, KaVo, Germany) using 50-pm white aluminum oxide blasting compound (Ivoclar Vivadent, Liechtenstein). The sandblasted discs were then sonicated (Branson 5200, Branson Ultrasonics, Danbury, CT, USA) twice in ultrapure water for 5 minutes each to remove any remnants of blasting compound. The discs were degreased with acetone for 15 minutes, then acid etched with 30% nitric acid for 30 minutes. The discs were then rinsed with ultrapure water and stored in 70% ethanol. Before use in transfection experiments, discs were sonicated twice for 15 min in ultrapure water, then rinsed with 70% ethanol, transferred to a biosafety cabinet, then rinsed thrice with sterile DNAse/RNAse-free water (ThermoFisher Scientific). The washed discs were fitted into 3D printed snaps, which were then slid into 3D printed bases glued to a microplate lid. The discs, snaps, and lid were all disinfected within the biosafety cabinet with ultraviolet light at 300 pW/cm² for 20 minutes. Polyplex solution was pipetted onto the surface of the discs and spread across the entire surface of the disc using a pipette tip. The suspension lid and discs covered with polyplex solution were covered with an inverted 24-well tissue culture microplate (DOT Scientific), and the polyplex solution was frozen on the discs in a -80°C freezer. The discs were then lyophilized (FreeZone 4.5 -105, Labconco).

Transfection efficiency assessment from titanium discs:

Cells were seeded in collagen-coated (collagen I, rat tail, Sigma Aldrich) 24-well tissue culture plates (DOT Scientific) 24 hours prior to transfection at a seeding density of 50,000 cells per well (for HEK293T cells) or 80,000 cells per well (primary gingival fibroblasts and keratinocytes). For each well, existing medium was replaced with 1.2 ml of complete DMEM lacking FBS. Coated discs attached to the suspension lid were removed from the lyophilizer and quickly placed on the plate containing the cells, resulting in the coated surface of the titanium discs being 4 mm from the bottom of the well. The plate with suspension lid and discs suspended above cells was gently shaken horizontally in two directions to ensure the medium contacted the entire surface of the discs. The plate was incubated under standard culture conditions for 4 hours, after which the suspension lid and discs were removed, and the medium in each well was replaced with 1 ml complete DMEM containing 10% (v/v) FBS. Forty-eight hours post- transfection, cells were trypsinized with 200 pl of 0.25% trypsin-EDTA (ThermoFisher Scientific), and 1 ml complete DMEM medium was added to each well to neutralize the trypsin. The detached cells were then suspended via pipetting. The resulting cell suspensions were transferred to 1 ml tubes and analyzed with a FACScan (Beckton Dickinson, Franklin Lakes, NJ, USA) flow cytometer equipped with a 15 mW of 488-nm excitation laser. Forward scatter, side scatter, and green fluorescence (FL1, 560-nm filter) parameters were measured. Cell debris was excluded through analysis with FlowJo software. A fluorescence threshold based on the negative control was created, and the percentage of cells fluorescing above the threshold was determined for each sample. Fluorescence microscopy (EVOS FL, ThermoFisher Scientific) was used to qualitatively assess transfection efficiency after transfection with polyplexes containing EGFP pDNA delivered from titanium discs.

Transfection efficiency after lyophilization in sucralose and long-term storage:

Polyplex solutions prepared as described above and containing 2% sucralose were placed into cryotubes, frozen at -80°C, and lyophilized (FreeZone 4.5 -105, Labconco). HEK 293T cells were seeded into collagen-coated (collagen I, rat tail, Sigma Aldrich) 24-well plates at a seeding density of 75,000 cells per well 24 hours prior to transfection. Polyplexes were either prepared the day of the experiment (as described above) or were resuspended from lyophilized polyplex solutions prepared earlier (24 hours to 2 years prior to the day of the experiment, also prepared as described above). Cells were treated with 2.5 pg of pDNA in polyplexes, then incubated for 48 hours. Transfection efficiency was assessed with flow cytometry and fluorescence microscopy, as described above.

Adsorption of polyplexes to CPC powder:

Polyplex solutions were prepared as described above and CPC prints were pulverized in a bead mill. The resulting CPC powder was mixed with 1 mL of polyplex solution for 5 minutes. The resulting mixture and 1 mL of polyplex solution not containing CPC powder were then centrifuged for 30 seconds (10,000g). HEK 293T cells were seeded onto collagen-coated (collagen I, rat tail, Sigma Aldrich) 24-well plates at a seeding density of 50,000 cells per well 24 hours prior to transfection. Cells were treated with non-centrifuged polyplex solution (Polyplexes), the supernatant of centrifuged polyplex solution (Poly. Supernatant), the supernatant of centrifuged CPC/polyplex mixture (CPC/Poly Supernatant), or the resuspended pellet from the centrifuged CPC/polyplex mixture (CPC/Poly Mixture), then incubated for 72 hours. Transfection efficiency was assessed with flow cytometry and fluorescence microscopy, as described above.

Gene -activation of CPC discs and meshes:

CPC discs were prepared as described above. For gene-activation of CPC discs with a lyophilized coating, CPC discs were incubated in complete DMEM for 24 hours, then polyplex solutions containing 2% sucralose were prepared as described above, then polyplex solutions containing 1.389 pg of pDNA were pipetted onto the discs. Polyplex-coated discs were then frozen at -80°C and lyophilized (FreeZone 4.5 -105, Labconco). For gene-activation of CPC discs via adsorption, CPC discs were incubated in complete DMEM for 24 hours, then polyplex solutions were prepared as described above, then incubated in polyplex solution (9 discs per 1.5 mL of polyplex solution) for 1 hour, after which the discs were immediately placed in wells for cell seeding to avoid polyplex inactivation due to drying. For gene-activation of CPC meshes, meshes were placed into 96- well plates containing lOOuLs of 4% (w/v) sucralose in pure water, then exposed to vacuum in a vacuum desiccator to remove air trapped in the nanotexture of the CPC. The meshes were then removed from the wells containing sucralose solution, placed in new 24-well plates, frozen at -80°C, then lyophilized (FreeZone 4.5 -105, Labconco). Lastly, 50 pL of polyplex solution (containing 5 pg of pDNA in polyplexes, prepared as described above) without sucralose was added to each mesh, after which the meshes were lyophilized again (FreeZone 4.5 -105, Labconco).

Transfection efficiency assessment from CPC:

HEK 293 T cells were seeded onto gene-activated CPC prints at a density based on the type of CPC print (disc or mesh). CPC discs were placed in 48-well plates and seeded with 37,500 cells per well. CPC meshes were placed in 12-well plates and seeded with 150,000 cells per well. Cells were incubated (72 hours for CPC discs, 96 hours for CPC meshes). Transfection efficiency was assessed with flow cytometry and fluorescence microscopy, as described above.

Statistical analysis:

All statistical analysis was performed with Prism (version 9, GraphPad, San Diego, CA, USA). If statistical comparisons were made between more than two groups within a dataset, the datasets were assessed with QQ plots, residual normality tests, and homoscedasticity tests to check whether assumptions of equal variance and normal distribution were violated. If a dataset failed the assumption of normal distribution, the Kruskal -Wallis test with Dunn’s multiple comparisons testing was used for statistical analysis. If a dataset passed the assumption of normality but failed the assumption of equal variance, then the Brown-Forsythe and Welch test with Dunnett’s T3 multiple comparisons testing was used for statistical analysis. If a dataset passed both the assumption of normal distribution and the assumption of equal variance, then ordinary one-way ANOVA with Tukey’s multiple comparisons testing was used for statistical analysis. The statistical tests used for each dataset are indicated in the figure legend.

The goals were to enhance the surface and mechanical properties of CPC scaffolds then gene-activate the resulting scaffolds. The first goal was achieved by showing that CPC scaffolds hardened with the VS hardening method had superior mechanical strength and improved osteogenic potential. The second goal was also accomplished, with our data demonstrating successful transfection of cells seeded onto gene-activated CPC discs and 3D meshes albeit with a different mode of transfection (surface-mediated transfection) than we expected. In the process of gene-activating surfaces and scaffolds, we also demonstrated for the first time that sucralose can perform as a lyoprotectant for lyophilization of PEI- pDNA polyplexes and can preserve them long-term. Our in vitro data suggests that the gene-activated scaffolds we developed should yield strong bone regeneration in vivo. This is supported by results from the literature that show highly nanotextured surfaces,⁴⁴’⁴⁵’⁸¹’⁹⁷-" gene-activation,^{65 112} and materials that locally induce a slightly alkaline pH^{108 109} all enhance bone regeneration.

The scaffold preparation method described herein includes secondary processing (wet-sanding, drilling) to give the scaffold the desired morphology. This is in contradiction with one of the main assumed benefits of 3D printing: being able to print a structure that replicates a specific 3D shape without additional processing. We decided to use the standard 3D object processing performed by our slicing software (which produces prints with occluded exterior pores) and then remove the occlusions after hardening. However, the printing parameters may be employed to prevent these occlusions, potentially by having the printhead break away from each printed strand before starting a new one. While we anticipated immediate release of polyplexes from our CPC 3D meshes based on the short incubation time before freezing and observations from our titanium and CPC discs (Fig. 18A), all the polyplexes were adsorbed to the surface of the mesh. This may actually be helpful with reducing off-target effects from gene-activated scaffolds. One of the concerns with using BMP-2 protein in bone tissue engineering is the fact that the protein can lead to off-target effects in regions local to the scaffold, most notably ectopic bone formation.⁶⁹ Polyplexes encoding BMP-2 immediately released into solution after implantation could have a greater chance that they could transfect cells in local tissue and potentially induce similar ectopic bone formation. However, it was demonstrated herein that the polyplexes loaded onto the scaffolds are incapable of transfecting cells not adhered to the surface of the scaffold, and as such our gene-activated scaffolds would likely be unable to transfect cells in local tissues.

Example 4:

Calcium phosphate cement (CPC) implants were prepared that were highly textured at the nanoscale (Fig- 11) The untreated CPC implants appeared hydrophobic upon exposure to water (Fig. 21A and 21C). This was surprising because the implants were made of calcium phosphate, which is made of polar ions, making it almost by definition a hydrophilic material and thus it was not expected to have exhibited hydrophobic properties. However, due to the nanotextured surface, the water-implant interface had entered a Cassie-Baxter state (Fig. 21E).

The Cassie-Baxter state is a liquid-surface interface state wherein air is trapped at the interface of a textured material and a liquid. This state leads to a hydrophilic material exhibiting a hydrophobic properties because of the entrapment of air in the surface’s texture. The Cassie-Baxter state has been intentionally induced in some materials via nanotexturing to create blood- repellant surfaces (i.e. non-wettable surfaces) that reduce the likelihood of blood clots forming upon exposure to blood. These types of surfaces have been investigated for use in medical devices where it is desirable to reduce blood clotting.

In contrast, wettability by blood is highly desirable for bone implants because high wettability facilitates improved cell-material interactions that improves osseointegration of the implant. Our implant treatment method is a process where a porous implant is combined with a solution containing a lyoprotectant, then the combination is degassed, then lyophilized (shown in schematic of Fig. 22). When the treated porous CPC implants were exposed to water, the water was absorbed readily and completely filled the interior of the implant (Figs. 21B and 21D). Fig. 23 depicts the treated porous CPC implant continuously absorbing water for a recorded period of 2.67 seconds.

Histological analysis of implants that were fixed in rabbit femurs showed differences in the amount of cell contact with the implant material (Fig. 24). Specifically, there are more cells and more bone formation on the surface of the implant (contact osteogenesis) in the center of the implant in the CPC implants treated with lyoprotectant-containing solution. This is due to the treatment enabling water/aqueous fluids (i.e. blood) to penetrate the center of the implant and make direct contact with the implant surface without being impeded by the formation of a Cassie-Baxter state at the fluid/implant interface. By filling void spaces within the CPC surface with lyoprotectant-containing solution then removing the water via lyophilization, system was created that can conduct water into the voids of the porous CPC implant’s nanostructured surface and bypass the Cassie-Baxter state. This method is not limited to application in implants made from CPC and is applicable to similar implants made from other materials, like titanium, polymers, or other minerals. The lyoprotectant solution used in this method also can be made from a variety of lyoprotectants (sucrose, sucralose, trehalose, etc.), and can also contain microparticles, gene delivery vectors, and/or other therapeutics.

References

1. Langer, R. & Vacanti, J. P. Tissue Engineering. Science 260, 920-926 (1993).

2. Laird, N. Z., Acri, T. M., Tingle, K. & Salem, A. K. Gene- and RNAi- activated scaffolds for bone tissue engineering: Current progress and future directions. Advanced Drug Delivery Reviews 174, 613-627 (2021).

3. Yan, W.-C. et al. 3D bioprinting of skin tissue: From pre-processing to final product evaluation. Advanced Drug Delivery Reviews 132, 270-295 (2018).

4. Snyder, Y. & Jana, S. Strategies for development of decellularized heart valve scaffolds for tissue engineering. Biomaterials 121675 (2022) doi: 10.1016/j. biomaterials.2022.121675. 5. Bhatia, S. N., Underhill, G. H., Zaret, K. S. & Fox, I. J. Cell and tissue engineering for liver disease. Science Translational Medicine 6, 245sr2- 245sr2 (2014).

6. Ferreira, A. M., Gentile, P., Chiono, V. & Ciardelli, G. Collagen for bone tissue regeneration. Acta Biomaterialia 8, 3191-3200 (2012).

7. Akkineni, A. R. et al. 3D plotting of growth factor loaded calcium phosphate cement scaffolds. Acta Biomaterialia 27, 264-274 (2015).

8. Bharadwaz, A. & Jayasuriya, A. C. Recent trends in the application of widely used natural and synthetic polymer nanocomposites in bone tissue regeneration. Mater Sci Eng C Mater Biol Appl 110, 110698 (2020).

9. Commissioner, O. of the. FDA Approves StrataGraft for the Treatment of Adults with Thermal Burns. FDA https://www.fda.gov/news-events/press- announcements/fda-approves-stratagraft-treatment-adults-thermal-burns (2021).

10. FDA approves Integra Omnigraft Dermal Regeneration Matrix to treat diabetic foot ulcers. FDA https://www.fda.gov/news-events/press- announcements/fda-approves-integra-omni graft-dermal-regeneration-matrix- treat-diabetic-foot-ulcers (2018).

11. Medtronic. Bone Grafting - Spine and Trauma Surgery - Infuse Bone Graft, https://www.medtronic.com/us-en/healthcare- professionals/products/spinal-orthopaedic/bone-grafting/infuse-bone- graft.html.

12. Commissioner, O. of the. FDA approves first autologous cellularized scaffold for the repair of cartilage defects of the knee. FDA https://www.fda.gov/news-events/press-announcements/fda-approves-first- autologous-cellularized-scaffold-repair-cartilage-defects-knee (2020).

13. De La Vega, R. E. et al. Efficient healing of large osseous segmental defects using optimized chemically modified messenger RNA encoding BMP -2. Science Advances 8, eabl6242 (2022).

14. Raftery, R. M. et al. Delivery of the improved BMP-2-Advanced plasmid DNA within a gene-activated scaffold accelerates mesenchymal stem cell osteogenesis and critical size defect repair. Journal of Controlled Release 283, 20-31 (2018). 15. Elangovan, S. et al. Chemically modified RNA activated matrices enhance bone regeneration. Journal of Controlled Release 218, 22-28 (2015).

16. Korn, P. et al. 3D Printing of Bone Grafts for Cleft Alveolar Osteoplasty

- In vivo Evaluation in a Preclinical Model. Front Bioeng Biotechnol 8, 217 (2020).

17. Raymond, Y. et al. 3D printing with star-shaped strands: A new approach to enhance in vivo bone regeneration. Biomaterials Advances 137, 212807 (2022).

18. Do, A.-V., Khorsand, B., Geary, S. M. & Salem, A. K. 3D Printing of Scaffolds for Tissue Regeneration Applications. Adv Healthc Mater 4, 1742- 1762 (2015).

19. 3D Printing Processes. Wikipedia (2022).

20. Landhuis, E. The Definition of Gene Therapy Has Changed. Nature (2021 ) doi : 10.1038/d41586-021 -02736-8.

21. Sherkow, J. S., Zettler, P. J. & Greely, H. T. Is it ‘gene therapy’? J Law Biosci 5, 786-793 (2018).

22. Research, C. for B. E. and. What is Gene Therapy? FDA (2020).

23. Kay, M. A. State-of-the-art gene-based therapies: the road ahead. Nat Rev Genet 12, 316-328 (2011).

24. Porto, E. M., Komor, A. C., Slaymaker, I. M. & Yeo, G. W. Base editing: advances and therapeutic opportunities. Nat Rev Drug Discov 19, 839-859 (2020).

25. Anzalone, A. V., Koblan, L. W. & Liu, D. R. Genome editing with CRISPR-Cas nucleases, base editors, transposases and prime editors. Nat Biotechnol 38, 824-844 (2020).

26. Mali, P. et al. RNA-Guided Human Genome Engineering via Cas9. Science 339, 823-826 (2013).

27. Jinek, M. et al. A Programmable Dual-RNA-Guided DNA Endonuclease in Adaptive Bacterial Immunity. Science 337, 816-821 (2012).

28. Baden, L. R. et al. Efficacy and Safety of the mRNA-1273 SARS-CoV-2 Vaccine. New England Journal of Medicine 384, 403-416 (2021).

29. Polack, F. P. et al. Safety and Efficacy of the BNT162b2 mRNA Covid-

19 Vaccine. New England Journal of Medicine 383, 2603-2615 (2020). . Fact Check-mRNA vaccines are distinct from gene therapy, which alters recipient’s genes. Reuters (2021). . Cohen, J. Pfizer-BioNTech And Moderna mRNA Vaccines Are Not Gene Therapies. Forbes https://www.forbes.eom/sites/joshuacohen/2022/01/04/pfizer-biontech-and- moderna-mma-vaccines-are-not-gene-therapies/. . Boussif, O. et al. A versatile vector for gene and oligonucleotide transfer into cells in culture and in vivo: polyethylenimine. Proc Natl Acad Sci USA 92, 7297-7301 (1995). . Wang, Z. et al. Detection of integration of plasmid DNA into host genomic DNA following intramuscular injection and electroporation. Gene Ther 11, 711-721 (2004). . Nauth, A., Schemitsch, E., Norris, B., Nollin, Z. & Watson, J. T. Critical- Size Bone Defects: Is There a Consensus for Diagnosis and Treatment?

Journal of Orthopaedic Trauma 32, S7 (2018). . Keating, J. F., Simpson, A. H. R. W. & Robinson, C. M. The management of fractures with bone loss. The Journal of Bone and Joint Surgery. British volume 87-B, 142-150 (2005). . Panetta, N. J. et al. Tissue Engineering in Cleft Palate and Other Congenital Malformations. Pediatric Research 63, 545-551 (2008). . Schemitsch, E. H. Size Matters: Defining Critical in Bone Defect Size! J Orthop Trauma 31 Suppl 5, S20-S22 (2017). . Roddy, E., DeBaun, M. R., Daoud-Gray, A., Yang, Y. P. & Gardner, M.

J. Treatment of critical-sized bone defects: clinical and tissue engineering perspectives. Eur J Orthop Surg Traumatol 28, 351-362 (2018). . Myeroff, C. & Archdeacon, M. Autogenous Bone Graft: Donor Sites and Techniques. The Journal of Bone & Joint Surgery 93, 2227-2236 (2011).. Agarwal, R. & Garcia, A. J. Biomaterial strategies for engineering implants for enhanced osseointegration and bone repair. Advanced Drug Delivery Reviews 94, 53-62 (2015). . Shaw, N. et al. Regenerative Medicine Approaches for the Treatment of Pediatric Physeal Injuries. Tissue Eng Part B Rev 24, 85-97 (2018). . Chanchareonsook, N., Junker, R., Jongpaiboonkit, L. & Jansen, J. A.

Tissue-Engineered Mandibular Bone Reconstruction for Continuity Defects: A Systematic Approach to the Literature. Tissue Eng Part B Rev 20, 147-162 (2014). . Boyan, B. D., Hummert, T. W Dean, D. D. & Schwartz, Z. Role of material surfaces in regulating bone and cartilage cell response. Biomaterials 17, 137-146 (1996). . Olivares-Navarrete, R. et al. Osteoblast lineage cells can discriminate microscale topographic features on titanium-aluminum-vanadium surfaces. Ann Biomed Eng 42, 2551-2561 (2014). . Lotz, E. M., Olivares-Navarrete, R., Berner, S., Boyan, B. D. & Schwartz, Z. Osteogenic response of human MSCs and osteoblasts to hydrophilic and hydrophobic nanostructured titanium implant surfaces. J Biomed Mater Res A 104, 3137-3148 (2016). . Barba, A. et al. Osteoinduction by Foamed and 3D-Printed Calcium Phosphate Scaffolds: Effect of Nanostructure and Pore Architecture. ACS AppL Mater. Interfaces 9, 41722-41736 (2017). . Park, S., Kim, G., Jeon, Y. C., Koh, Y. & Kim, W. 3D polycaprolactone scaffolds with controlled pore structure using a rapid prototyping system. J Mater Sci Mater Med 20, 229-234 (2009). . Lu, J. X. et al. Role of interconnections in porous bioceramics on bone recolonization in vitro and in vivo. J Mater Sci Mater Med 10, 111-120 (1999). . Melchels, F. P. W. et al. Effects of the architecture of tissue engineering scaffolds on cell seeding and culturing. Acta Biomater 6, 4208-4217 (2010).. Kuboki, Y., Jin, Q. & Takita, H. Geometry of Carriers Controlling Phenotypic Expression in BMP-Induced Osteogenesis and Chondrogenesis. JBJS K3, S105 (2001). . Lee, J. W ., Ahn, G., Kim, J. Y. & Cho, D.-W. Evaluating cell proliferation based on internal pore size and 3D scaffold architecture fabricated using solid freeform fabrication technology. J Mater Sci Mater Med 21, 3195-3205 (2010). . Lim, T. C., Chian, K. S. & Leong, K. F. Cryogenic prototyping of chitosan scaffolds with controlled micro and macro architecture and their effect on in vivo neo-vascularization and cellular infiltration. J Biomed Mater Res A 94, 1303-1311 (2010). 53. Mygind, T. et al. Mesenchymal stem cell ingrowth and differentiation on coralline hydroxyapatite scaffolds. Biomaterials 28, 1036-1047 (2007).

54. Matai, I., Kaur, G., Seyedsalehi, A., McClinton, A. & Laurencin, C. T. Progress in 3D bioprinting technology for tissue/organ regenerative engineering. Biomaterials 226, 119536 (2020).

55. Han, H. H. et al. Reconstruction of Complex Maxillary Defects Using Patient-specific 3D-printed Biodegradable Scaffolds. Plast Reconstr Surg Glob Open 6, (2018).

56. He, S., Fang, J., Zhong, C., Wang, M. & Ren, F. Spatiotemporal Delivery of pBMP2 and pVEGF by a Core-Sheath Structured Fiber-Hydrogel Gene- Activated Matrix Loaded with Peptide-Modified Nanoparticles for Critical- Sized Bone Defect Repair. Advanced Healthcare Materials n/a, 2201096.

57. Moncal, K. K. et al. Controlled Co-delivery of pPDGF-B and pBMP-2 from intraoperatively bioprinted bone constructs improves the repair of calvarial defects in rats. Biomaterials 281, 121333 (2022).

58. Rowland, C. R. et al. Regulation of decellularized tissue remodeling via scaffold-mediated lentiviral delivery in anatomically-shaped osteochondral constructs. Biomaterials 177, 161-175 (2018).

59. Khorsand, B. et al. A Multi-Functional Implant Induces Bone Formation in a Diabetic Model. Adv Healthc Mater 9, e2000770 (2020).

60. Khorsand, B., Elangovan, S., Hong, L., Kormann, M. S. D. & Salem, A.

K. A bioactive collagen membrane that enhances bone regeneration. J Biomed Mater Res B Appl Biomater 107, 1824-1832 (2019).

61. Khorsand, B. et al. A Comparative Study of the Bone Regenerative Effect of Chemically Modified RNA Encoding BMP-2 or BMP-9. AABS J 19, 438-446 (2017).

62. Khorsand, B. et al. Regeneration of bone using nanoplex delivery of FGF-2 and BMP-2 genes in diaphyseal long bone radial defects in a diabetic rabbit model. Journal of Controlled Release 248, 53-59 (2017).

63. Walsh, D. P. et al. Gene activated scaffolds incorporating star-shaped polypeptide-pDNA nanomedicines accelerate bone tissue regeneration in vivo. Biomater. Sci. 9, 4984-4999 (2021).

64. Lackington, W. A. et al. Non-viral Gene Delivery of Interleukin- 1 Receptor Antagonist Using Collagen-Hydroxyapatite Scaffold Protects Rat BM-MSCs From IL-ip-Mediated Inhibition of Osteogenesis. Front Bioeng Biotechnol , 582012 (2020).

65. Raftery, R. M. et al. Highly versatile cell-penetrating peptide loaded scaffold for efficient and localised gene delivery to multiple cell types: From development to application in tissue engineering. Biomaterials 216, 119277 (2019).

66. Gonzalez-Fernandez, T., Tierney, E. G., Cunniffe, G. M., O’Brien, F. J.

& Kelly, D. J. Gene Delivery of TGF-P3 and BMP2 in an MSC-Laden Alginate Hydrogel for Articular Cartilage and Endochondral Bone Tissue Engineering. Tissue Eng Part A 22, 776-787 (2016).

67. Premarket Approval (PMA) P050053 (Infuse Bone Graft). https://www.accessdata.fda.gov/scripts/cdrh/cfdocs/cfpma/pma.cfm?id=P050 053.

68. Rengachary, S. S. Bone morphogenetic proteins: basic concepts. Neurosurgical Focus 13, 1-6 (2002).

69. James, A. W. et al. A Review of the Clinical Side Effects of Bone Morphogenetic Protein-2. Tissue Eng Part B Rev 22, 284-297 (2016).

70. Woo, E. J. Adverse events reported after the use of recombinant human bone morphogenetic protein 2. J Oral Maxillofac Surg 70, 765-767 (2012).

71. Carragee, E. J., Hurwitz, E. L. & Weiner, B. K. A critical review of recombinant human bone morphogenetic protein-2 trials in spinal surgery: emerging safety concerns and lessons learned. Spine J 11, 471-491 (2011).

72. Lee, K., Silva, E. A. & Mooney, D. J. Growth factor delivery-based tissue engineering: general approaches and a review of recent developments. J R Soc Interface 8, 153-170 (2011).

73. Mroz, T. E., Wang, J. C., Hashimoto, R. & Norvell, D. C. Complications Related to Osteobiologics Use in Spine Surgery: A Systematic Review. Spine 35, S86 (2010).

74. Tannoury, C. A. & An, H. S. Complications with the use of bone morphogenetic protein 2 (BMP-2) in spine surgery. Spine J 14, 552-559 (2014).

75. Zara, J. N. et al. High Doses of Bone Morphogenetic Protein 2 Induce Structurally Abnormal Bone and Inflammation In Vivo. Tissue Eng Part A 17, 1389-1399 (2011). . Kolk, A. et al. Comparative analysis of bone regeneration behavior using recombinant human BMP-2 versus plasmid DNA of BMP-2. J Biomed Mater Res A 107, 163-173 (2019). . Raper, S. E. et al. Fatal systemic inflammatory response syndrome in a ornithine transcarbamylase deficient patient following adenoviral gene transfer. Molecular Genetics and Metabolism 80, 148-158 (2003). . Egermann, M. et al. Effect of BMP-2 gene transfer on bone healing in sheep. Gene Ther 13, 1290-1299 (2006). . Wu, M., Chen, G. & Li, Y.-P. TGF-P and BMP signaling in osteoblast, skeletal development, and bone formation, homeostasis and disease. Bone Research 4, 1-21 (2016). . Xu, H. H. et al. Calcium phosphate cements for bone engineering and their biological properties. Bone Res 5, 17056 (2017). . Barba, A. et al. Impact of Biomimicry in the Design of Osteoinductive Bone Substitutes: Nanoscale Matters. ACS Appl. Mater. Interfaces 11, 8818— 8830 (2019). . LeGeros, R. Z. Calcium Phosphate-Based Osteoinductive Materials. ('hem. Rev. 108, 4742-4753 (2008). . Ginebra, M.-P., Canal, C., Espanol, M., Pastorino, D. & Montufar, E. B. Calcium phosphate cements as drug delivery materials. Advanced Drug Delivery Reviews 64, 1090-1110 (2012). . Yousefi, A.-M. A review of calcium phosphate cements and acrylic bone cements as injectable materials for bone repair and implant fixation. Journal of Applied Biomaterials & Functional Materials 17, 2280800019872594 (2019). . Heinemann, S., Rbssler, S., Lemm, M., Ruhnow, M. & Nies, B. Properties of injectable ready -to-use calcium phosphate cement based on water-immiscible liquid. Acta Biomaterialia 9, 6199-6207 (2013). . Lode, A. et al. Fabrication of porous scaffolds by three-dimensional plotting of a pasty calcium phosphate bone cement under mild conditions. Journal of Tissue Engineering and Regenerative Medicine 8, 682-693 (2014).. Zhu, H., Guo, D., Qi, W. & Xu, K. Development of Sr-incorporated biphasic calcium phosphate bone cement. Biomed. Mater. 12, 015016 (2017). 88. Kim, Y., Uyama, E., Sekine, K., Kawano, F. & Hamada, K. Effects of pol oxamer additives on strength, injectability, and shape stability of betatricalcium phosphate cement modified using ball-milling. Journal of the Mechanical Behavior of Biomedical Materials 130, 105182 (2022).

89. Koh, I., Lopez, A., Pinar, A. B., Helgason, B. & Ferguson, S. J. The effect of water on the mechanical properties of soluble and insoluble ceramic cements. Journal of the Mechanical Behavior of Biomedical Materials 51, 50-60 (2015).

90. Sadowska, J.-M., Guillem-Marti, J., Montufar, E. B., Espanol, M. & Ginebra, M.-P. Biomimetic Versus Sintered Calcium Phosphates: The In Vitro Behavior of Osteoblasts and Mesenchymal Stem Cells. Tissue Engineering Part A 23, 1297-1309 (2017).

91. Liu, H. et al. Reinforcement of injectable calcium phosphate cement by gelatinized starches. Journal of Biomedical Materials Research Part B : Applied Biomaterials 104, 615-625 (2016).

92. Komath, M. & Varma, H. K. Development of a fully injectable calcium phosphate cement for orthopedic and dental applications. Bull Mater Sci 26, 415-422 (2003).

93. Ahlfeld, T. et al. 3D Plotted Biphasic Bone Scaffolds for Growth Factor Delivery: Biological Characterization In Vitro and In Vivo. Adv Healthc Mater 8, el801512 (2019).

94. Kilian, D. et al. 3D Bioprinting of osteochondral tissue substitutes - in vitro-chondrogenesis in multi-layered mineralized constructs. Sci Rep 10, 8277 (2020).

95. Bagnol, R. et al. Coaxial micro-extrusion of a calcium phosphate ink with aqueous solvents improves printing stability, structure fidelity and mechanical properties. Acta Biomaterialia 125, 322-332 (2021).

96. Chow, L. C. & Takagi, S. A Natural Bone Cement — A Laboratory Novelty Led to the Development of Revolutionary New Biomaterials. J Res Natl Inst Stand Technol 106, 1029-1033 (2001).

97. Wang, M. et al. Cold atmospheric plasma (CAP) surface nanomodified 3D printed polylactic acid (PLA) scaffolds for bone regeneration. Acta Biomater 46, 256-265 (2016). 98. Nowicki, M. A., Castro, N. J., Plesniak, M. W. & Zhang, L. G. 3D printing of novel osteochondral scaffolds with graded microstructure. Nanotechnology 27, 414001 (2016).

99. Hesaraki, S., Nazarian, H., Pourbaghi-Masouleh, M. & Borhan, S. Comparative study of mesenchymal stem cells osteogenic differentiation on low-temperature biomineralized nanocrystalline carbonated hydroxyapatite and sintered hydroxyapatite. Journal of Biomedical Materials Research Part B: Applied Biomaterials 102, 108-118 (2014).

100. INNOTERE Biomaterial I Plotter Paste-CPC. innotere-neu https://www.innotere.de/plotter-paste-cpc.

101. Ahlfeld, T. et al. Design and Fabrication of Complex Scaffolds for Bone Defect Healing: Combined 3D Plotting of a Calcium Phosphate Cement and a Growth Factor-Loaded Hydrogel. Ann Biomed Eng 45, 224-236 (2017).

102. Ahlfeld, T. et al. Bioprinting of mineralized constructs utilizing multichannel plotting of a self-setting calcium phosphate cement and a cellladen bioink. Biofabrication 10, 045002 (2018).

103. Acri, T. et al. Non-viral gene delivery embedded in biomimetically mineralized matrices for bone tissue engineering. Tissue Eng Part A (2020) doi: 10.1089/ten.TEA.2020.0206.

104. Klimek, K. et al. “False” cytotoxicity of ions-adsorbing hydroxyapatite — Corrected method of cytotoxicity evaluation for ceramics of high specific surface area. Materials Science and Engineering: C 65, 70-79 (2016).

105. Sadowska, J. M. et al. In vitro response of mesenchymal stem cells to biomimetic hydroxyapatite substrates: A new strategy to assess the effect of ion exchange.

-332 (2018).

106. Gustavsson, J., Ginebra, M. P., Engel, E. & Planell, J. Ion reactivity of calcium -defi ci ent hydroxyapatite in standard cell culture media. Acta Biomaterialia 7, 4242-4252 (2011).

107. Rollin-Martinet, S., Navrotsky, A., Champion, E., Grossin, D. & Drouet, C. Thermodynamic basis for evolution of apatite in calcified tissues.

American Mineralogist 98, 2037-2045 (2013).

108. Shen, Y. et al. Interfacial pH: a critical factor for osteoporotic bone regeneration. Langmuir 27, 2701-2708 (2011). 109. Shen, Y. et al. Bone regeneration: importance of local pH — strontium- doped borosilicate scaffold. J. Mater. Chem. 22, 8662-8670 (2012).

110. Kohn, D. H., Sarmadi, M., Helman, J. I. & Krebsbach, P. H. Effects of pH on human bone marrow stromal cells in vitro: Implications for tissue engineering of bone. Journal of Biomedical Materials Research 60, 292-299 (2002).

111. D’Mello, S., Elangovan, S. & Salem, A. K. FGF2 gene activated matrices promote proliferation of bone marrow stromal cells. Arch Oral Biol 60, 1742-1749 (2015).

112. Elangovan, S. et al. The enhancement of bone regeneration by gene activated matrix encoding for platelet derived growth factor. Biomaterials 35, 737-747 (2014).

113. Hall, A., Lachelt, U., Bartek, J., Wagner, E. & Moghimi, S. M. Polyplex Evolution: Understanding Biology, Optimizing Performance. Molecular Therapy 25, 1476-1490 (2017).

114. Benjaminsen, R. V., Mattebjerg, M. A., Henriksen, J. R., Moghimi, S. M. & Andresen, T. L. The Possible “Proton Sponge ” Effect of Polyethylenimine (PEI) Does Not Include Change in Lysosomal pH. Molecular Therapy 21, 149-157 (2013).

115. Cohen, R. N., van der Aa, M. A. E. M., Macaraeg, N., Lee, A. P. & Szoka, F. C. Quantification of plasmid DNA copies in the nucleus after lipoplex and polyplex transfection. Journal of Controlled Release 135, 166- 174 (2009).

116. Grandinetti, G. & Reineke, T. M. Exploring the Mechanism of Plasmid DNA Nuclear Internalization with Polymer-Based Vehicles. Mol. Pharmaceutics 9, 2256-2267 (2012).

117. Laird, N. Z. et al. A proof of concept gene-activated titanium surface for oral implantology applications. Journal of Tissue Engineering and Regenerative Medicine 14, 622-632 (2020).

118. Gervasi, V. et al. Parenteral protein formulations: An overview of approved products within the European Union. European Journal of Pharmaceutics and Biopharmaceutics 131, 8-24 (2018).

119. Jorgensen, L., Hostrup, S., Moeller, E. H. & Grohganz, H. Recent trends in stabilising peptides and proteins in pharmaceutical formulation - considerations in the choice of excipients. Expert Opinion on Drug Delivery 6, 1219-1230 (2009).

120. Jakubovics, N. S., Goodman, S. D., Mashburn- Warren, L., Stafford, G. P. & Cieplik, F. The dental plaque biofilm matrix. Periodontology 2000 86, 32- 56 (2021).

121. Souza, J. G. S. et al. Effect of sucrose on biofilm formed in situ on titanium material. J. Periodontol. 90, 141-148 (2019).

122. Mensink, M. A., Frij link, H. W ., van der Voort Maarschalk, K. & Hinrichs, W. L. J. How sugars protect proteins in the solid state and during drying (review): Mechanisms of stabilization in relation to stress conditions. European Journal of Pharmaceutics and Biopharmaceutics 114, 288-295 (2017).

123. Roberts, A., Renwick, A. G., Sims, J. & Snodin, D. J. Sucralose metabolism and pharmacokinetics in man. Food and Chemical Toxicology 38, 31-41 (2000).

124. Omran, A., Ahearn, G., Bowers, D., Swenson, J. & Coughlin, C. Metabolic effects of sucralose on environmental bacteria. J Toxicol 2013, 372986 (2013).

125. What Is Splenda (Sucralose) And How Is It Made? | Splenda FAQ.

Splenda® https://www.splenda.com/faqs/what-is-splenda-sucralose-and-how- is-it-made/.

126. Malkawi, W. I. et al. Application of lyophilized gene-delivery formulations to dental implant surfaces: non-cariogenic lyoprotectant preserves transfection activity of polyplexes long-term. J Pharm Sci (2022).

127. Meng-Lund, H. et al. Exploring the chemical space for freeze-drying excipients. Int J Pharm 566, 254-263 (2019).

128. Holm, T. P., Meng-Lund, H., Rantanen, J., Jorgensen, L. & Grohganz, H. Screening of novel excipients for freeze-dried protein formulations. Eur J Pharm Biopharm 160, 55-64 (2021).

129. Imberti, S., McLain, S. E., Rhys, N. H., Bruni, F. & Ricci, M. A. Role of Water in Sucrose, Lactose, and Sucralose Taste: The Sweeter, The Wetter? ACS Omega 4, 22392-22398 (2019). 130. Madry, H., Venkatesan, J. K., Carballo-Pedrares, N., Rey-Rico, A. & Cucchiarini, M. Scaffold-Mediated Gene Delivery for Osteochondral Repair. Pharmaceutics 12, E930 (2020).

131. Raftery, R. M., Tierney, E. G., Curtin, C. M., Cryan, S.-A. & O’Brien, F. J. Development of a gene-activated scaffold platform for tissue engineering applications using chitosan-pDNA nanoparticles on collagen-based scaffolds. Journal of Controlled Release 210, 84-94 (2015).

132. D’Mello, S., Salem, A. K., Hong, L. & Elangovan, S. Characterization and evaluation of the efficacy of cationic complex mediated plasmid DNA delivery in human embryonic palatal mesenchyme cells. J Tissue Eng Regen Med ia, 927-937 (2016).

133. Atluri, K., Lee, J., Seabold, D., Elangovan, S. & Salem, A. K. Gene- Activated Titanium Surfaces Promote In Vitro Osteogenesis. IntJ Oral Maxillofac Implants 32, e83-e96 (2017).

134. Choi, S., Yu, X., Jongpaiboonkit, L., Hollister, S. J. & Murphy, W. L. Inorganic coatings for optimized non-viral transfection of stem cells. Sci Rep 3, 1567 (2013).

135. Olivares-Navarrete, R. et al. Osteoblast lineage cells can discriminate microscale topographic features on titanium-aluminum-vanadium surfaces. Ann Biomed Eng 42, 2551-2561 (2014).

136. Lotz, E. M., Olivares-Navarrete, R., Berner, S., Boyan, B. D. & Schwartz, Z. Osteogenic response of human MSCs and osteoblasts to hydrophilic and hydrophobic nanostructured titanium implant surfaces. J Biomed Mater Res A 104, 3137-3148 (2016).

137. Wang, M. et al. Cold atmospheric plasma (CAP) surface nanomodified 3D printed polylactic acid (PLA) scaffolds for bone regeneration. Acta Biomater 46, 256-265 (2016).

138. Nowicki, M. A., Castro, N. J., Plesniak, M. W. & Zhang, L. G. 3D printing of novel osteochondral scaffolds with graded microstructure. Nanotechnology 27, 414001 (2016).

139. Barba, A. et al. Impact of Biomimicry in the Design of Osteoinductive Bone Substitutes: Nanoscale Matters. ACS Appl. Mater. Interfaces 11, 8818-8830 (2019). 140. Hesaraki, S., Nazarian, H., Pourbaghi-Masouleh, M. & Borhan, S. Comparative study of mesenchymal stem cells osteogenic differentiation on low- temperature biomineralized nanocrystalline carbonated hydroxyapatite and sintered hydroxyapatite. Journal of Biomedical Materials Research Part B: Applied Biomaterials 102, 108-118 (2014).

141. Liu, J.-L., Feng, X.-Q., Wang, G. & Yu, S.-W. Mechanisms of superhydrophobicity on hydrophilic substrates. J. Phys.: Condens. Matter 19, 356002 (2007).

142. Hosono, E., Fujihara, S., Honma, I. & Zhou, H. Superhydrophobic Perpendicular Nanopin Film by the Bottom-Up Process. J. Am. Chem. Soc. 127, 13458-13459 (2005).

143. Herminghaus, S. Roughness-induced non-wetting. EPL 52, 165 (2000).

144. Feng, L. et al. Creation of a Superhydrophobic Surface from an Amphiphilic Polymer. Angewandte Chemie International Edition 42, 800-802 (2003).

145. Zhu, M., Zuo, W., Yu, H., Yang, W. & Chen, Y. Superhydrophobic surface directly created by electrospinning based on hydrophilic material. J Mater Sci 41, 3793-3797 (2006).

146. Jokinen, V., Kankuri, E., Hoshian, S., Franssila, S. & Ras, R. H. A. Superhydrophobic Blood-Repellent Surfaces. Advanced Materials 30, 1705104 (2018).

147. Gittens, R. A. et al. A review on the wettability of dental implant surfaces II: Biological and clinical aspects. Acta Biomaterialia 10, 2907-2918 (2014).

All patents and publications referenced or mentioned herein are indicative of the levels of skill of those skilled in the art to which the invention pertains, and each such referenced patent or publication is hereby specifically incorporated by reference to the same extent as if it had been incorporated by reference in its entirety individually or set forth herein in its entirety. Applicants reserve the right to physically incorporate into this specification any and all materials and information from any such cited patents or publications.

The following statements are intended to describe and summarize various embodiments of the invention according to the foregoing description in the specification. Statements:

1. A method to prepare a porous scaffold with hydrophilic pores, comprising: treating the porous scaffold with a composition comprising one or more lyoprotectants; degassing the treated porous scaffold; freezing the porous scaffold; and lyophilizing the scaffold.

2. The method of statement 1 wherein the porous scaffold is immersed in a solution comprising the composition.

3. The method of statement 1 or 2 wherein the one or more lyoprotectants comprise sucrose, trehalose or sucralose.

4. The method of statement 1, 2 or 3 wherein the composition comprises water.

5. The method of any one of statements 1 to 4 wherein the lyoprotectant is about 0.5% w/v to about 20% w/v in the composition.

6. The method of any one of statements 1 to 4 wherein the lyoprotectant is about 2.0% w/v to about 10% w/v in the composition.

7. The method of any one of statements 1 to 4 wherein the lyoprotectant is about 3.0% w/v to about 5% w/v in the composition.

8. The method of any one of statements 1 to 7 wherein the scaffold comprises calcium phosphate.

9. The method of any one of statements 1 to 8 wherein the scaffold comprises microstructures or nanostructures.

10. A product produced by the method of any one of statements 1 to 9. 11. A method, comprising; contacting the product of statement 10 and a composition comprising nucleic acid.

12. The method of statement 11 wherein the nucleic acid comprises DNA.

13. The method of statement 11 wherein the nucleic acid comprises RNA including chemically modified RNA.

14. The method of statement 11 wherein the nucleic acid is a plasmid.

15. The method of any one of statements 11 to 14 wherein the nucleic acid encodes a gene product that enhances bone formation or regeneration.

16. The method of any one of statements 11 to 14 wherein the composition comprises a polymer.

17. The method of statement 16 wherein the polymer is a natural polymer.

18. The method of statement 16 wherein the polymer is a synthetic polymer.

19. The method of statement 18 wherein the polymer comprises PEI or

PLGA.

20. The method of statement 1, wherein degassing the treated porous scaffold comprises applying a vacuum, an ultrasound, or a combination thereof.

The specific methods, devices and compositions described herein are representative of preferred embodiments and are exemplary and not intended as limitations on the scope of the invention. Other objects, aspects, and embodiments will occur to those skilled in the art upon consideration of this specification, and are encompassed within the spirit of the invention as defined by the scope of the claims. It will be readily apparent to one skilled in the art that varying substitutions and modifications can be made to the invention disclosed herein without departing from the scope and spirit of the invention.

The invention illustratively described herein suitably can be practiced in the absence of any element or elements, or limitation or limitations, which is not specifically disclosed herein as essential. The methods and processes illustratively described herein suitably can be practiced in differing orders of steps, and the methods and processes are not necessarily restricted to the orders of steps indicated herein or in the claims.

Under no circumstances can the patent be interpreted to be limited to the specific examples or embodiments or methods specifically disclosed herein. Under no circumstances can the patent be interpreted to be limited by any statement made by any Examiner or any other official or employee of the Patent and Trademark Office unless such statement is specifically and without qualification or reservation expressly adopted in a responsive writing by Applicants.

The terms and expressions that have been employed are used as terms of description and not of limitation, and there is no intent in the use of such terms and expressions to exclude any equivalent of the features shown and described or portions thereof, but it is recognized that various modifications are possible within the scope of the invention as claimed. Thus, it will be understood that although the present invention has been specifically disclosed by preferred embodiments and optional features, modification and variation of the concepts herein disclosed can be resorted to by those skilled in the art, and that such modifications and variations are considered to be within the scope of this invention as defined by the appended claims and statements of the invention.

The invention has been described broadly and generically herein. Each of the narrower species and subgeneric groupings falling within the generic disclosure also form part of the invention. This includes the generic description of the invention with a proviso or negative limitation removing any subject matter from the genus, regardless of whether or not the excised material is specifically recited herein. In addition, where features or aspects of the invention are described in terms of Markush groups, those skilled in the art will recognize that the invention is also thereby described in terms of any individual member or subgroup of members of the Markush group.

Claims

WHAT IS CLAIMED IS:

1. A method to prepare a porous scaffold with hydrophilic pores, comprising: treating the porous scaffold with a composition comprising one or more lyoprotectants; degassing the treated porous scaffold; freezing the porous scaffold; and lyophilizing the scaffold.

2. The method of claim 1 wherein the porous scaffold is immersed in a solution comprising the composition.

3. The method of claim 1 or 2 wherein the one or more lyoprotectants comprise sucrose, trehalose or sucralose.

4. The method of claim 1, 2 or 3 wherein the composition comprises water.

5. The method of any one of claims 1 to 4 wherein the lyoprotectant is about 0.5% w/v to about 20% w/v in the composition.

6. The method of any one of claims 1 to 4 wherein the lyoprotectant is about 2.0% w/v to about 10% w/v in the composition.

7. The method of any one of claims 1 to 4 wherein the lyoprotectant is about 3.0% w/v to about 5% w/v in the composition.

8. The method of any one of claims 1 to 7 wherein the scaffold comprises calcium phosphate.

9. The method of any one of claims 1 to 8 wherein the scaffold comprises microstructures or nanostructures.

10. A product produced by the method of any one of claims 1 to 9.

11. A method, comprising; contacting the product of claim 10 and a composition comprising nucleic acid.

12. The method of claim 11 wherein the nucleic acid comprises DNA.

13. The method of claim 11 wherein the nucleic acid comprises RNA including chemically modified RNA.

14. The method of claim 11 wherein the nucleic acid is a plasmid.

15. The method of any one of claims 11 to 14 wherein the nucleic acid encodes a gene product that enhances bone formation or regeneration.

16. The method of any one of claims 11 to 14 wherein the composition comprises a polymer.

17. The method of claim 16 wherein the polymer is a natural polymer.

18. The method of claim 16 wherein the polymer is a synthetic polymer.

19. The method of claim 18 wherein the polymer comprises PEI or PLGA.

20. The method of claim 1, wherein degassing the treated porous scaffold comprises applying a vacuum, an ultrasound, or a combination thereof.