Micro-Nanofibrillar Polycaprolactone Scaffolds as Translatable

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Micro-Nanofibrillar Polycaprolactone Scaffolds as Translatable Osteoconductive Grafts for the Treatment of Musculoskeletal Defects Without Infection Paria Ghannadian, James Moxley, Mirian Michelle Machado de Paula, Anderson O. Lobo, and Thomas J. Webster ACS Appl. Bio Mater., Just Accepted Manuscript • DOI: 10.1021/acsabm.8b00453 • Publication Date (Web): 12 Oct 2018 Downloaded from http://pubs.acs.org on October 15, 2018

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Figure 1. Demonstration of the three fabrication techniques implemented for the synthesize of PCL scaffolds: ES, RJS, and AB. (A.) Schematic of the ES extrusion process; (B.) Schematic of the RJS extrusion process; (C.) Schematic of the AB extrusion process. ACS Paragon Plus Environment

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Figure 2. Demonstration of fiber dimensions and pore sizing for ES, RJS, and AB scaffolds. (A.) SEM imaging of an ES 12 % PCL (w/v) microfiber at 10,000x magnification; (B.) SEM imaging of a RJS 12 % PCL (w/v) microfiber at 10,000x magnification; (C.) SEM imaging of two pairs of AB 12 % PCL (w/v) nanofibers at 10,000x magnification; (D.) Average fiber diameters for the PCL scaffolds, with standard deviation values, determined through SEM imaging (at 700x magnification) and subsequent ImageJ processing; (E.) Average pore diameters for the PCL scaffolds, with standard deviation values, determined through SEM imaging (at 700 x magnification) and subsequent ImageJ processing. Scale bars are 1 µm in length. ACS Paragon Plus Environment

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200 160

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Figure 3. Contact angle measurements for distilled water droplets applied to the PCL scaffolds at ambient temperature and pressure. Note that *p < 0.05, ***p < 0.001, and n = 2. 80 Elastic Modulus (MPa)

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Figure 4. Elastic modulus values determined for the PCL scaffolds via tensile tests. Note that *p < 0.05, **p < 0.01, ***p < 0.001, and n = 3.


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600 Fracture Strain (%)

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Figure 5. Fracture strain values determined for the PCL scaffolds via tensile tests. Note that *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001, and n = 3.


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Figure 6. Cell adhesion assay for the PCL scaffolds, assessed at 4 hours following the initial seeding event. Note that ***p < 0.001, ****p < 0.0001, and n = 6.


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Figure 7. Metabolic activity assay for the PCL scaffolds, assessed at 24 hours following the initial seeding event. Note that *p < 0.05, **p < 0.01, ***p < 0.001, and n = 6.


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Figure 8. Cell proliferation assay for the PCL scaffolds, assessed at 168 hours and 336 hours following the initial seeding event. Note that **p < 0.01, ***p < 0.001, ****p < 0.0001, and n = 6.


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Figure 9. Demonstration of critically-point dried hFOB cells within the PCL scaffolds at 168 hours following the initial seeding event. (A.) SEM imaging of hFOB cells within an ES 12 % PCL (w/v) scaffold, at 2,000x magnification. Observe that the cells have wrapped around individual microfibers, and have contacts extending through multiple planes in the scaffold. (B.) SEM imaging of hFOB cells within a RJS 12 % PCL (w/v) scaffold, at 2,000x magnification. Observe that the cells have spread over a large surface area, and that indentations from the microfibrous scaffolding are apparent along the cellular sheet. (C.) SEM imaging of hFOB cells within an AB 12 % PCL (w/v) scaffold, at 2,000x magnification. Observe that the cells have colonized the intraporous space of the nanofibrous scaffolding, through interstrand adhesions. Scale bars are 20 µm in length. ACS Paragon Plus Environment

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(A.) COL1A1

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3 2 1 0 8% ES

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Figure 10. Gene expression studies for the PCL scaffolds, assessed at 336 hours following the initial seeding event. Relative gene expression is demonstrated for: (A.) COL1A1, (B.) COL1A2, (C.) ALPL, (D.) SPP1, and (E.) BGLAP. Relative gene expression was calculated by the ΔΔCT method, with GAPDH serving as the reference housekeeping gene. Note that n = 3.


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Figure 11. Alkaline phosphatase bioactivity within the PCL scaffolds, assessed at 168 hours and 336 hours following the initial seeding event. Note that *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001, and n = 6.



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Figure 12. Calcium deposition within the PCL scaffolds, assessed at 168 hours and 336 hours following the initial seeding event. Note that ****p < 0.0001, and n = 6. ACS Paragon Plus Environment

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(A.) S. aureus *

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Figure 13. Colony forming unit assay for (A.) gram-positive S. aureus and (B.) gram-negative P. aeruginosa bacterial strains within the PCL scaffolds, assessed at 24 hours following the initial inoculation event. Note that *p < 0.05, **p < 0.01, ***p < 0.001, and n = 9. ACS Paragon Plus Environment



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1 Title: Micro-Nanofibrillar Polycaprolactone Scaffolds as Translatable Osteoconductive Grafts for the Treatment of Musculoskeletal Defects Without Infection Paria Ghannadian1,2,† James Walter Moxley, Junior1,2,† Mirian Michelle Machado de Paula2,5,6 Anderson de Oliveira Lobo2,3,4,5,6,7 Thomas Jay Webster1,2,§ 1Chemical

Engineering Department, Northeastern University, Boston, Massachusetts, 02115, USA.

2Nanomedicine

Laboratory, Northeastern University, Boston, Massachusetts, 02115, USA.

3Department

of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Boston, Massachusetts, 02115, USA. 4Harvard

– MIT Health Sciences and Technology, Massachusetts Institute of Technology, Cambridge, Massachusetts, 02139, USA. 5Department

of Biomedical Engineering, Universidade do Vale do Paraíba, São José dos Campos, São Paulo, 12244 – 000, Brazil. 6Laboratory

of Biomedical Nanotechnology, Universidade do Vale do Paraíba, São José dos Campos, São Paulo, 12244 – 000, Brazil. 7Interdisciplinary

Laboratory for Advanced Materials, Universidade Federal do Piauí, Teresina, Piauí, 64049 – 550, Brazil. †Authors

Contributed Equally to this Work.

§Corresponding

Author: Dr. Thomas J. Webster, 313 Snell Engineering Center, 360 Huntington Avenue, Boston, MA, 02115, USA. [email protected]


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2 Table of Contents  Title ………………………………………………………………………………………………………………………………………….. 1  Authors and Affiliations ……………………………………………………………………………………………………………. 1  Abstract ……………………………………………………………………………………………………………………………………. 3  Keywords ……………………………………………………………………………………………………………………………….... 3  Introduction ……………………………………………………………………………………………………………………………… 4  Materials  Synthesis ……………………………………………………………………………………………………………………… 6  Materials Characterization and Mechanical Testing …………………………………………………….. 6  Osteoblast Viability Evaluation …………………………………………………………………………………….. 6  Osteoblast Phenotype Evaluation ………………………………………………………………………………… 7  Bacterial Viability Evaluation ………………………………………………………………………………………… 8  Methods  Synthesis ……………………………………………………………………………………………………………………… 8  Materials Characterization and Mechanical Testing …………………………………………………… 12  Osteoblast Viability Evaluation …………………………………………………………………………………… 12  Osteoblast Phenotype Evaluation ………………………………………………………………………………. 14  Bacterial Viability Evaluation ……………………………………………………………………………………… 17  Results and Discussion  Materials Characterization and Mechanical Testing …………………………………………………… 18  Osteoblast Viability Evaluation …………………………………………………………………………………… 23  Osteoblast Phenotype Evaluation ………………………………………………………………………………. 28  Bacterial Viability Evaluation ……………………………………………………………………………………… 34  Conclusions …………………………………………………………………………………………………………………………….. 36  Acknowledgements ………………………………………………………………………………………………………………… 39  Institution Identification …………………………………………………………………………………………………………. 39  Funding Sources ……………………………………………………………………………………………………………………… 39  References ……………………………………………………………………………………………………………………………… 40



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3 Abstract: The treatment of musculoskeletal defects is currently limited by the tissue-regenerative materials available to orthopedic surgeons: autologous bone grafts only have a finite amount of harvestable material within a given patient, while allografts are prone to severe immunological complications and host rejection. With this motivation, the production of poly(ε – caprolactone) (PCL) scaffolds as synthetic, biomimetic biomaterials was investigated, with specific focus on potential orthopedic translation. PCL scaffolds were produced through three different fabrication techniques: Electrospinning (ES), Rotary Jet Spinning (RJS), and Airbrush (AB). ES and RJS were observed to produce microfibrillar scaffolds, while all AB products were nanofibrous. Osteoblast viability within the PCL scaffolds, as well as the osteogenic phenotype, were assessed in vitro through a combination of adherence, metabolic activity, proliferation, gene expression, alkaline phosphatase bioactivity, and calcium deposition assays. While the polymeric scaffolds induced slight reductions in initial osteoblast adhesion and metabolic activity, seeded cells were able to proliferate and demonstrate the bone formation phenotype. AB products demonstrated reduced bacterial surface colonization when inoculated with both gram-positive (Staphylococcus aureus) and gram-negative (Pseudomonas aeruginosa) bacterial strains, in comparison to the microfibrous ES and RJS products, without any small-molecule antibiotics, antimicrobial peptides, or reactive nanomaterials included during scaffold synthesis.

Keywords: Orthopedic Biomaterial; Antibacterial Nanomaterial; Osteoconductive Tissue Engineering Scaffold; Osteogenic Phenotype; Polycaprolactone Micro-Nanofibers.


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4 Introduction: Orthopedic surgeons are currently limited by materials available for the treatment of large-scale musculoskeletal damage stemming from traumatic injury, congenital birth defects, and previous surgical excision. Among current-generation treatments, autologous bone grafts are frequently cited as the de facto gold standard, but this approach is limited by the amount of harvestable material for transplantation as well as significant donor site morbidity.1-3 As an alternative to autologous grafts, allografts are still ultimately limited by tissue availability within donor banks, and are generally more susceptible to severe immunological complications (delaying osseointegration and promoting host-transplant rejection) than autografts.4,5 Synthetic, biomimetic biomaterials are currently being explored as next-generation alternatives for highly-translatable, orthopedic implants which are capable of regenerating the structure and function of native bone tissue, without the constraints of availability and immunoclearance. Fibrillar scaffolds are often studied as potential orthopedic biomaterials, as they are largely analogous to endogenous extracellular matrices found within a given patient.6,7 From a design perspective, the minimal functional requirement is that these structures are osteoconductive, providing a substrate for native osteoblasts to infiltrate and direct localized osteogenesis.8 These scaffolds are microporous, which promotes facile cell migration, as well as providing ample free space for the diffusion of nutrients, wastes, and secretory signaling species. More advanced designs are also functionalized with osteoinductive ligands, which promote the development of native mesenchymal and osteoprogenitor cells to fully-differentiated osteoblasts within the treatment site.9-11 As osteoblasts will direct the formation of new bone tissue, through the production and biomineralization of the osteoid, an ideal scaffolding substrate should be both biodegradable and bioresorbable.12-14 Poly(ε – caprolactone) (PCL) was explored in this study for its utilization in the fabrication of microscale and nanoscale fibrillar scaffolds which can be clinically implemented as next-generation, orthopedic implants. PCL is a synthetic, biodegradable polymer, and its degradation products (caproic



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5 acid) may be either renally excreted or directly consumed as a cellular metabolite.15 PCL has been approved by the United States Food and Drug Administration (FDA) for specific in vivo applications, and has been thoroughly studied for implementation in surgical sutures as well as drug delivery devices.16 Although the polymer has a lower tensile stress, and thus does not effectively emulate the mechanical properties of mineralized musculoskeletal tissue, its high elasticity may prove advantageous for molding and manipulation during surgical implantation. PCL-based biomaterials often demonstrate strong hydrophobic character, but this can be modulated through the application of plasma treatment to induce intrastrand breaks and introduce polar functional groups along the modified surfaces.17,18 In addition to increasing hydrophilicity, plasma treatment presents an avenue for introducing chemical signaling within PCL scaffolds, through the conjugation of biomolecular species (such as osteoinductive ligands, adhesion peptides, and antimicrobial peptides) to said functional groups.19,20 PCL demonstrates high processability, and can incorporate a diverse array of nanomaterials (such as carbon nanotubes and hydroxyapatite nanoparticulates) which can further modulate the mechanical properties and degradation rate of the resultant composite material.21-24 In the course of this study, unmodified PCL scaffolds (which have neither received plasma treatment nor incorporated nanomaterials) were produced through three fabrication methods: Electrospinning (ES), Rotary Jet Spinning (RJS), and Airbrush (AB). The osteoconductive potential of these polymeric scaffolds was assessed in vitro, with specific focus on osteoblast viability, osteoid production, and biomineralization. In addition to osteogenic studies, the relative susceptibility of these polymeric scaffolds to gram-positive and gram-negative bacterial infections was also assessed in vitro. Bacterial infections of biomaterials is a pressing global challenge due to the increasing frequency of occurrence of antibiotic resistant (ABR) bacterial strains, specifically multidrug resistant (MDR), extensively drug resistant (XDR), and totally drug resistant (TDR) phenotypes. Bacterial infections with ABR stains have been projected to induce greater fatalities than all cancer types combined by the year 2050.25


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6 The focus of this study was specifically placed on assessing the potential clinical significance of unmodified PCL scaffolds as osteoconductive implants. By eliminating any post-manufacture processing of the scaffolding, as well as forgoing the incorporation of any nanomaterials or biological additives within the matrix design, clear steps were taken to reduce both production time and costs, while simultaneously reducing the opportunity for batch-to-batch variability. While these engineering choices are clearly attractive from the standpoint of potential industrial scale-up, they also provide a more direct pathway for rapid regulatory approval and subsequent clinical implementation. Materials: Synthesis – Micro-nanoscale fibrillar scaffolds were synthesized through three different solvent-evaporation methods, utilizing PCL (number average molecular weight of 80,000 Daltons, 440744, Sigma-Aldrich) and 1,1,1,3,3,3 – Hexafluoro – 2 – propanol (HFIP, 003409, Oakwood Chemical) as the polymer and solvent species. A Double Spinner (Inovenso), MR – 115 Mini – Rotary (AWT), and Model G23 Airbrush (Master Airbrush) were utilized for ES, RJS, and AB production, respectively. Materials Characterization and Mechanical Testing – Materials characterization and mechanical testing were conducted on empty polymeric scaffolds with a S – 4800 SEM (Hitachi), RSA – G2 Solids Analyzer (TA Instruments), and Phoenix 150 Contact Angle Analyzer (Surface Electro Optics). These scaffolds are referred to as empty as assessments were conducted prior to either seeding with mammalian cells or inoculating with bacteria. Osteoblast Viability Evaluation – Eukaryotic cellular response upon seeding within the polymeric scaffolds was assessed with the use of the hFOB 1.19 cell line (CRL – 11372, ATCC). The hFOB cells were cultured in Gibco DMEM/F – 12 No Phenol Red (21 – 041 – 025, Thermo Fisher), supplemented with 10 % FBS (30 – 2020, ATCC) and a 1 % penicillin – streptomycin solution (15140122, Thermo Fisher). Cellular adhesion, metabolic activity, and



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7 proliferation were assessed through the implementation of the CellTiter 96 Aqueous One Solution Cell Proliferation Assay (G3581, Promega). Polymeric scaffolds seeded with hFOB cells were fixed with a solution of 25 % glutaraldehyde in water (G5882, Sigma – Aldrich) and 4 % paraformaldehyde (PFA) in phosphate buffered saline (PBS) (J61899, Alfa Aesar), critical-point dried with a Samdri – PVT – 3D (Tousimis), and analyzed with the S – 4800 SEM. Osteoblast Phenotype Evaluation – The expression of multiple gene sequences related to osteogenesis was assessed with a TaqMan Gene Expression Cells to CT Kit (AM1728, Thermo Fisher Scientific), utilizing primers for COL1A1 (4331182 – Hs00164004_m1, Thermo Fisher Scientific), COL1A2 (4331182 – Hs01028956, Thermo Fisher Scientific), ALPL (4331182 – Hs01029144_m1, Thermo Fisher Scientific), SPP1 (4331182 – Hs00959010_m1, Thermo Fisher Scientific), BGLAP (4331182 – Hs00609452_g1, Thermo Fisher Scientific), and GAPDH (Hs02786624_g1, Thermo Fisher Scientific) as a reference control. Gene expression studies implemented an Arktik Thermal Cycler (Thermo Scientific) for reverse transcription (RT), and a QuantStudio 6 Flex Real – Time PCR System (Applied Biosystems) for quantitative polymerase chain reaction (qPCR) analysis. The progression of osteogenesis within the seeded polymeric scaffolds was more directly assessed with the implementation of a QuantiChrom Alkaline Phosphatase Assay Kit (DALP – 250, BioAssay Systems) and BioAssay Systems Calcium Assay Kit (50 – 489 – 211, Thermo Fisher Scientific). Triton X – 100 (T8787, Sigma Aldrich) was utilized as a cell lytic reagent during the assessment of alkaline phosphatase bioactivity, while hydrochloric acid (H1758, Sigma Aldrich) was utilized as a dissolving solution for calcium deposition evaluation. All spectrophotometric readings conducted during the evaluation of osteoblast viability and phenotype were completed with a SpectraMax M3 Multi-Mode Microplate Reader (Molecular Devices),


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8 utilizing SoftMax Pro 6 Software (Molecular Devices). Similarly, all experimental steps requiring a set period of agitation were completed with an Innova 2000 (New Brunswick) platform shaker. Bacterial Viability Evaluation – Prokaryotic cellular response upon inoculation within the polymeric scaffolds was assessed with the use of Staphylococcus aureus (12600, ATCC) and Pseudomonas aeruginosa (35984, ATCC) cell lines. The S. aureus and P. aeruginosa cells were cultured in tryptic soy broth (TSB) (22092, Sigma – Aldrich), with colony counting conducted on plated agar (A1296, Sigma – Aldrich).

Methods: Synthesis – PCL pellets were dissolved in HFIP to effectively produce polymeric solutions with 8 %, 12 %, and 17 % PCL concentration (w/v). These polymeric solutions were subsequently extruded through an orifice, and directed at a collecting surface. As the polymeric solution passed through the air, between the orifice and the collecting surface, the solvent species was removed via evaporation due to its high volatility. This resulted in the deposition of solid, nonwoven mats of micro-nanoscale PCL fibrils along the surface of the collector. The extrusion process was conducted through three potential fabrication techniques: ES, RJS, and AB. ES utilizes an applied electric field in order generate a difference in electric potential between the collecting surface and the polymer solution held within the orifice. It is this difference in electric potential which will ultimately induce and drive extrusion, and as such ES may be considered an electrochemical extrusion process (Figure 1, A). Once the difference in electric potential is sufficient to overcome the surface tension of the polymeric solution, this solution will develop a characteristic Taylor cone deformation and be propelled out of the orifice as a dynamic jet stream. In this experiment, a Double Spinner with a gauge size of 23 (orifice internal diameter of 0.33 mm) was utilized for ES fabrication. A



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9 voltage range of 5 – 7 kV was applied to the system, with the ejected polymeric solution collected on a rotating mandrel over a 6-hour period. Extrusion was conducted at ambient temperature and pressure, with the observed humidity ranging from 30 % – 45 % in the local environment. Polymeric solutions with 8 %, 12 %, and 17 % PCL concentration (w/v) were capable of producing microfibril scaffolds through the ES process. RJS utilizes centrifugal force in order to drive the polymeric solution out of a rotating orifice, and as such it may be considered a physical extrusion process (Figure 1, B). In this experiment, an RJS apparatus was constructed by mounting a rotatable nozzle, with a gauge size of 24 (orifice internal diameter of 0.311 mm), on a MR – 115 Mini – Rotary. The nozzle was rotated perpendicular to its rotary central axis at a speed of 3,000 rpm, with the ejected polymeric solution collected on a circular arrangement of stationary mandrels over a 1-hour period. Extrusion was conducted at ambient temperature and pressure, with the observed humidity ranging from 30 % – 45 % in the local environment. Polymeric solutions with 8 %, 12 %, and 17 % PCL concentration (w/v) were capable of producing microfibril scaffolds through the RJS process. AB utilizes pneumatics in order to drive the polymeric solution out of an orifice, which may or may not be stationary depending on the fabrication procedure, and as such it may also be considered a physical extrusion process (Figure 1, C). In this experiment, a Model G23 Airbrush was utilized to propel compressed air, within a pressure range of 20 – 40 pounds per square inch (PSI), into a nozzle with a gauge size of 24. As the nozzle is continuously loaded with the polymeric solution, the stream of compressed air will induce continuous extrusion to a stationary, planar collector surface over a 2-hour period. Extrusion was conducted at ambient temperature and pressure, with the observed humidity ranging from 30 % – 45 % in the local environment. Polymeric solutions with 8 % and 12 % PCL concentration (w/v) were capable of producing nanofibril scaffolds through the AB process. Polymeric solutions with 17 % PCL concentration (w/v) proved too viscous for efficient extrusion through the airbrush system studied.


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12 Figure 1. Demonstration of the three fabrication techniques implemented for the synthesize of PCL scaffolds: ES, RJS, and AB. (A.) Schematic of the ES extrusion process; (B.) Schematic of the RJS extrusion process; (C.) Schematic of the AB extrusion process. Materials Characterization and Mechanical Testing – Materials characterization and mechanical testing were conducted on polymeric scaffolds that were neither seeded with mammalian cells nor inoculated with bacteria. The morphology of the micronanofibrils was initially observed with an S – 4800 SEM at 700X magnification. The average fiber diameter and pore sizing for the polymeric scaffolds was determined by processing the generated SEM images with ImageJ software. Contact angles were determined with a Phoenix 150 Contact Angle Analyzer, utilizing distilled water at ambient temperature and pressure. Images generated with this instrumentation were similarly processed with ImageJ software; however, in this instance the DropSnake plugin was specifically implemented. Tensile tests were conducted on rectangular samples of the polymeric scaffolds, specifically cut to have dimensions of 10.00 mm x 30.00 mm x 0.50 mm. These samples were fixed within a RSA – G2 Solids Analyzer using clamps provided by the instrumentation, with the gap control reduced to 1.00 mm in order to effectively remove all external tension from the samples prior to testing. Tensile tests generated a stress-strain curve for each of the polymeric scaffolds, which was subsequently utilized to calculate their elastic modulus and fracture strain. Osteoblast Viability Evaluation – In this experimental study, all work with mammalian cell cultures was based on the hFOB 1.19 cell lineage. The growth medium utilized was Gibco DMEM/F – 12 No Phenol Red, supplemented with 10 % FBS and a 1 % penicillin – streptomycin solution. Growth media was effectively replaced every 48 hours during the course of mammalian cell culturing.



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13 Osteoblast viability within the different polymeric scaffolds was assessed by means of an adhesion assay, metabolic activity assay, and proliferation assay. Each of these assessments utilized a variation of the standardized MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2Htetrazolium) assay protocol. Polymeric scaffolds were cut into square samples, with specific dimensions of 20.00 mm x 20.00 mm x 0.50 mm, and sterilized with two rounds of germicidal UVc radiation, each conducted for 20 minutes in duration. For the adhesion assay, the polymeric scaffolds were set into 24-well microplates and seeded with 10,000 hFOB cells in 1000 µL of growth media per well. Following an incubation period of 4 hours, at 34.0 degrees Celsius and 5.0 % atmospheric CO2, the scaffolds are transferred to new microplates with fresh growth media. A 200.0 µL aliquot of MTS was applied to each well, and these microplates were set into the incubator for an additional 4-hour development period prior to colorimetric assessment. Following the conclusion of the development period, spectrophotometric absorbance readings were conducted at a wavelength of 570 nm with the SpectraMax M3 Multi-Mode Microplate Reader, utilizing SoftMax Pro 6 Software. For the metabolic activity assay, the polymeric scaffolds were similarly set into 24-well microplates, but in this instance were seeded with 20,000 hFOB cells in 1000 µL of growth media per well. Following an incubation period of 24 hours, at 34.0 degrees Celsius and 5.0 % atmospheric CO2, a 200.0 µL aliquot of MTS was applied to each well and the microplates were returned to the incubator for an additional 4-hour development period prior to colorimetric assessment. Following the conclusion of the development period, spectrophotometric absorbance readings were conducted at a wavelength of 570 nm with the SpectraMax M3 Multi-Mode Microplate Reader, utilizing SoftMax Pro 6 Software. For the proliferation assay, the polymeric scaffolds were similarly set into 24-well microplates, but in this instance, they were seeded with 5,000 hFOB cells in 1000 µL of growth media per well. Following an incubation period of either 168 hours (7 days) or 336 hours (14 days), at 34.0 degrees Celsius


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14 and 5.0 % atmospheric CO2, a 200.0 µL aliquot of MTS was applied to each well and the microplates were returned to the incubator for an additional 4-hour development period prior to colorimetric assessment. Following the conclusion of the development period, spectrophotometric absorbance readings were conducted at a wavelength of 570 nm with the SpectraMax M3 Multi-Mode Microplate Reader, utilizing SoftMax Pro 6 Software. In order to effectively visualize osteoblasts cultured within the polymeric scaffolds, hFOB cells were fixed, critical-point dried, and imaged with a S – 4800 SEM. Polymeric scaffolds were set into 24-well microplates, and seeded with 5,000 hFOB cells in 1000 µL of growth media per well. Following an incubation period of 168 hours, at 34.0 degrees Celsius and 5.0 % atmospheric CO2, the growth media was removed from each well via aspiration and the polymeric scaffolds were washed with 1000 µL of PBS. A 25 % glutaraldehyde solution in water was diluted by a factor of ten, utilizing 4 % PFA in PBS as the diluent, to produce a fixative. After the wash solution is removed via aspiration, 1000 µL of the fixative was applied to the polymeric scaffolds and the microplates were incubated for 1 hour at room temperature. At the conclusion of this fixation stage, the fixative was removed via aspiration, and replaced with an initial 1000 µL aliquot of 30 % (v/v) ethanol in water. The polymeric scaffolds were incubated at room temperature for 10 minutes, at which point the initial solution was replaced with a second 1000 µL aliquot of 30 % (v/v) ethanol in water for an additional 10-minute incubation period. This process was repeated to progressively dehydrate the polymeric scaffolds, serially increasing the ethanol concentration to 50 % (v/v), 70 % (v/v), 80 % (v/v), 90 % (v/v), and finally 100 % (v/v). Once the PCL scaffolds were completely dehydrated, they were critical-point dried with a Samdri – PVT – 3D and imaged with a S – 4800 SEM at multiple magnifications. Osteoblast Phenotype Evaluation – The osteoblast phenotype, specifically bone formation functionality, was assessed by means of gene expression, alkaline phosphatase activity, and calcium deposition. While gene expression studies



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15 focused on the relative mRNA levels of genes associated with native osteogenesis, alkaline phosphatase activity and calcium deposition assay demonstrated the functionality of a specific osteogenic protein and overall biomineralization within the scaffolds, respectively. For the conduction of this study, polymeric scaffolds were once again cut into square samples, with specific dimensions of 20.00 mm x 20.00 mm x 0.50 mm, and sterilized with two rounds of germicidal UVc radiation, each conducted for 20 minutes in duration. The gene expression studies focused on several genes previously observed to be related to osteoblast-directed bone formation, specifically COL1A1 (expresses collagen, type 1, alpha 1 – a component of type I collagen), COL1A2 (expresses collagen, type 1, alpha 2 – a component of type I collagen), ALPL (expresses alkaline phosphatase, tissue non-specific isozyme), SPP1 (expresses osteopontin), and BGLAP (expresses osteocalcin). Gene expression for GAPDH (expresses glyceraldehyde 3-phosphate dehydrogenase) was also measured as a reference control. Polymeric scaffolds were set into 24-well microplates and seeded with 10,000 hFOB cells in 1000 µL of growth media. Following an incubation period of 336 hours, at 34.0 degrees Celsius and 5.0 % atmospheric CO2, hFOB cells residing within the polymeric scaffolds were effectively lysed and reverse transcribed to cDNA utilizing proprietary buffer solutions (Lysis Solution, Stop Solution, and RT Buffer) and enzymes (DNase I and RT Enzyme Mix) provided by the TaqMan Gene Expression Cells to CT Kit. cDNA prepared from the lysate solutions was analyzed via qPCR cycling within the QuantStudio 6 Flex Real – Time PCR System, again utilizing proprietary buffer solutions and enzymes (TaqMan Gene Expression Master Mix) provided by the TaqMan Gene Expression Cells to CT Kit, in addition to the select primers from the TaqMan Gene Expression Assays that are specific to the gene sequences identified above. For the alkaline phosphatase activity assay, the polymeric scaffolds were set into 24-well microplates and seeded with 5,000 hFOB cells in 1000 µL of growth media per well. Following an incubation period of either 168 hours or 336 hours, at 34.0 degrees Celsius and 5.0 % atmospheric CO2,


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16 the growth media was removed via aspiration and the polymeric scaffolds were transiently washed with PBS. Following aspiration of the PBS wash solution, 500.0 µL of lysis solution, consisting of 0.2 % Triton X – 100 (w/v) in distilled water, was applied to each scaffold. The 24-well microplates were immediately agitated at 125 rpm, for a duration of 20 minutes, on an Innova 2000 platform shaker held at ambient temperature and pressure. At the conclusion of the lysis stage, 50.0 µL of lysate from each scaffold was transferred to individual wells of a new 96-well microplate. At this juncture, the alkaline phosphatase reaction solution was produced by mixing 2.00 µL of 10 mM para-nitrophenylphosphate solution, 5.00 µL of 5 mM magnesium acetate solution, and 200.0 µL of proprietary alkaline phosphatase assay buffer, all provided by the QuantiChrom Alkaline Phosphatase Assay Kit. An aliquot of 150.0 µL of the alkaline phosphatase reaction solution was administered to the individual lysate solutions, which were subsequently homogenized by pipette trituration. Optical density measurements for the homogeneous reaction mixtures were conducted immediately at a wavelength of 405 nm with the SpectraMax M3 MultiMode Microplate Reader, utilizing SoftMax Pro 6 Software. Optical density measurements were repeated on the same mixtures after the reaction was allowed to proceed for 4 minutes, with the microplates incubated in the dark at ambient temperature and pressure for the duration of this period. For the calcium deposition assay, the polymeric scaffolds were similarly set into 24-well microplates, and seeded with 5,000 hFOB cells in 1000 µL of growth media per well. Following an incubation period of either 168 hours or 336 hours, at 34.0 degrees Celsius and 5.0 % atmospheric CO2, growth media was removed via aspiration and the polymeric scaffolds were washed three times with distilled water. Following aspiration of the final wash cycle, 500 µL of 0.6 M hydrochloric acid was administered to each scaffold. The 24-well microplates were immediately agitated at 125 rpm, for a duration of 4 hours, on an Innova 2000 platform shaker held at ambient temperature and pressure. At the conclusion of the dissolution stage, 500 µL of the resultant solution was extracted from each well and centrifuged individually at 12,000 rpm for 3 minutes at 4 degrees Celsius. Following centrifugation, 5.00



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17 µL of the supernatant solution was transferred to individual wells of a new 96-well microplate. At this juncture, the calcium deposition reaction solution was produced by mixing 100.0 µL of Reagent A (0.1 % – 2.0 % o-phthalic acid, w/v; 0.1 % – 1.0 % hydroxyquinoline, w/v) and 100.0 µL of Reagent B (0.1 % – 3.0 % sodium sulfite, w/v; 1.0 % – 20.0 % monoethanolamine, w/v), both supplied by the BioAssay Systems Calcium Assay Kit. An aliquot of 200.0 µL of calcium deposition reaction solution was administered to the individual supernatant solutions, which were subsequently homogenized by pipette trituration. Optical density measurements for the homogeneous reaction mixtures were conducted immediately at a wavelength of 612 nm with the SpectraMax M3 Multi-Mode Microplate Reader, utilizing SoftMax Pro 6 Software. Bacterial Viability Evaluation – Prokaryotic colonization and growth within the polymeric scaffolds was assessed with a variation of the standardized colony forming unit assay. S. aureus and P. aeruginosa were utilized in this experimental study, such that responses from both gram-positive and gram-negative bacterial strains were observed. For the conduction of this experimentation, polymeric scaffolds were once again cut into square samples, with specific dimensions of 20.00 mm x 20.00 mm x 0.50 mm, and sterilized with two rounds of germicidal UVc radiation, each conducted for 20 minutes in duration. A single colony of both S. aureus and P. aeruginosa were selected from their respective starter plates and were inoculated within separate conical tubes, each containing 5.00 mL of TSB. The bacterial solutions were both incubated under agitation at 120 rpm for 14 hours, utilizing an Innova 2000 platform shaker held at 37 degrees Celsius. At the conclusion of this primary incubation period, 100.0 µL of each bacterial culture was administered to separate wells of a 96-well microplate. Spectrophotometric absorbance measurements were conducted at a wavelength of 600 nm with the SpectraMax M3 MultiMode Microplate Reader, utilizing SoftMax Pro 6 Software, to determine that the bacterial cultures had a cell concentration of 1.2 x 109 cells/mL for S. aureus and 1.3 x 109 cells/mL for P. aeruginosa, respectively.


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18 Both bacterial cultures subsequently underwent serial dilutions in order to produce effective working concentrations of 1 x 104 cells/mL. Polymeric scaffolds were set into 24-well microplates, individually inoculated with 1000 µL of either S. aureus or P. aeruginosa working solution, and incubated for 24 hours at 37.0 degrees Celsius and 5.0 % atmospheric CO2. At the conclusion of this secondary incubation period, the bacterial cultures were removed via aspiration, and the polymeric scaffolds were subjected to three washing stages, each utilizing 1000 µL of PBS, in order to remove the subpopulation of loosely adherent bacteria. Immediately following the final washing stage, the polymeric scaffolds were set into individual conical tubes, each containing 1000 µL of PBS, and vortexed vigorously for 15 minutes to displace strongly adherent bacteria. After vortexing was complete, the polymeric scaffolds were removed from the conical tubes, and the bacterial suspensions were diluted to assessment solutions: the S. aureus suspensions were each diluted to assessment solutions with concentrations of 1 x 102 cells/mL, 1 x 103 cells/mL, and 1 x104 cells/mL; while the P. aeruginosa suspensions were each diluted to assessment solutions with concentrations of 1 x 104 cells/mL, 1 x 105 cells/mL, and 1 x 106 cells/mL. These assessment solutions were seeded onto agar plates as 10 µL aliquots, in triplicate, and incubated for 14 hours at 37.0 degrees Celsius and 5.0 % atmospheric CO2. At the conclusion of this final incubation period, colony counting was conducted on the seeded agar plates for both bacterial strains. Results and Discussion: Materials Characterization and Mechanical Testing – SEM imaging, and subsequent ImageJ processing, determined the average fiber dimensions and pore sizes for each of the polymeric scaffolds synthesized (Figure 2). PCL scaffolds fabricated via ES and RJS demonstrated fibrils within the micrometer range, while those produced via AB demonstrated fibrils on the order of hundreds of nanometers. Average pore sizes were largely comparable between the three fabrication techniques implemented within this study, and were on the order of several micrometers in diameter. A general trend of increasing average fiber diameter was observed with increasing PCL



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19 concentration, for all fabrication techniques. This trend was similarly observed for increasing average pore diameter with increasing PCL concentration. RJS demonstrated the greatest inherent variance in synthesis, with the largest standard deviation values for fiber and pore diameter across the range of PCL concentrations tested. Interestingly, ES demonstrated the greatest percent increase for both fiber diameter and pore diameter when moving from the lowest PCL concentration to the highest PCL concentration tested. (A.)

(B.)


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20 (C.)

(D.)



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21 (E.)

Figure 2. Demonstration of fiber dimensions and pore sizing for ES, RJS, and AB scaffolds. (A.) SEM imaging of an ES 12 % PCL (w/v) microfiber at 10,000x magnification; (B.) SEM imaging of a RJS 12 % PCL (w/v) microfiber at 10,000x magnification; (C.) SEM imaging of two pairs of AB 12 % PCL (w/v) nanofibers at 10,000x magnification; (D.) Average fiber diameters for the PCL scaffolds, with standard deviation values, determined through SEM imaging (at 700x magnification) and subsequent ImageJ processing; (E.) Average pore diameters for the PCL scaffolds, with standard deviation values, determined through SEM imaging (at 700 x magnification) and subsequent ImageJ processing. Scale bars are 1 µm in length. Contact angle measurements demonstrated that all polymeric scaffolds produced from the three fabrication techniques were, by definition, hydrophobic surfaces – although to varying degrees (Figure 3).26 A general trend of increasing hydrophobic character was observed with increasing PCL concentration in ES, RJS, and AB products. The contact angle measured for the lowest PCL concentration, 8 % PCL (w/v), was quite comparable between ES, RJS, and AB. Although the percent change in contact angle between the lowest PCL concentration and the highest PCL concentration was relatively small for ES and AB products, this value was quite pronounced for RJS. Perhaps most interesting is that the majority of this


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22 increase in contact angle, and thus hydrophobicity, was observed between the RJS 8 % PCL (w/v) and RJS 12 % PCL (w/v) scaffolds. The percent change between these two scaffolds was greater than that witnessed between the RJS 12 % PCL (w/v) and RJS 17 % PCL (w/v) scaffolds, by over an order of magnitude. Although all the polymeric scaffolds produced within the course of this study are clearly hydrophobic surfaces, in comparison to the hydrophilic, tissue-culture treated control surface, only the RJS 12 % PCL (w/v) and RJS 17 % PCL (w/v) scaffolds approach the superhydrophobic region.27,28 200 160

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Figure 3. Contact angle measurements for distilled water droplets applied to the PCL scaffolds at ambient temperature and pressure. Note that *p < 0.05, ***p < 0.001, and n = 2. Tensile tests demonstrated two significant mechanical properties for our polymeric scaffolds: elastic modulus and fracture strain. Both properties were observed to increase in value with increasing PCL concentration during scaffold synthesis. A material with a higher elastic modulus will be relatively stiff, and resistant to deformation under a given load, while a material with a lower elastic modulus will be comparably more flexible.29,30 With this understanding, the stiffest polymeric scaffold observed was AB 12 % PCL (w/v), while the most flexible was ES 8 % PCL (w/v) (Figure 4). The ES products demonstrated the greatest change in elastic modulus with increasing PCL concentration, and thus a greatest increase in stiffness, with an increase in value of over 3,000 %. Conversely, the RJS products demonstrated the smallest change in elastic modulus with increasing PCL concentration, with an overall increase of only 120.0 %. With regards to fracture strain, a material with a greater observed fracture strain will inherently be capable of greater resistance to deformation without onset cracking and fissure.31,32 ES products



23 demonstrated the highest fracture strain values upon tensile testing, followed by RJS products with moderate values, and AB products with considerably lower values (Figure 5). The polymeric scaffolds most resistant to cracking and fissure were ES 17 % PCL (w/v), while the polymeric scaffolds with the least resistance were AB 8 % PCL (w/v).

Elastic Modulus (MPa)

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Figure 4. Elastic modulus values determined for the PCL scaffolds via tensile tests. Note that *p < 0.05, **p < 0.01, ***p < 0.001, and n = 3. 600 Fracture Strain (%)

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Figure 5. Fracture strain values determined for the PCL scaffolds via tensile tests. Note that *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001, and n = 3. Osteoblast Viability Evaluation – With materials characterization and mechanical properties effectively determined for the PCL scaffolds, an assessment of hFOB adhesion, metabolism, and proliferation was utilized in order to evaluate osteoblast viability within these three-dimensional structures. The cell adhesion assay demonstrated that the hydrophobic, polymeric scaffolds had a significant reduction in adherent cells, at a time point 4 hours following the initial seeding event, when compared


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24 against a tissue-culture treated control surface (Figure 6). The average number of adherent cells measured for each of the polymeric scaffolds tested was below the lower standard deviation calculated for the control surface. The observed decrease in adherent cells in comparison to the control ranges from – 36. 1 % for the ES 17 % PCL (w/v) scaffold to – 67.0 % for the AB 12 % PCL (w/v) scaffold. This result was not surprising, as studies in the literature have demonstrated that the hydrophilicity, or hydrophobicity, of a given material will influence its intermolecular interactions with a cell line (whether eukaryotic or prokaryotic) upon attachment and early surface colonization.33-37 With regards to interactions with mammalian cell lineages, such as osteoblasts, it is the proteome of the plasma membrane which will dictate attachment; specifically, the expression of different integrin subunits, within the supramolecular adhesion complexes, may be of great significance.38,39 Previous studies focused on osteoblast attachment and proliferation rates have demonstrated improved culturing performance with materials of increasing hydrophilic nature.40-42 Through this lens, the observed reduction in initial hFOB adhesion was anticipated for the hydrophobic, PCL scaffolds. General trends in reduced adherence could not be immediately discerned between the different fabrication techniques or PCL concentrations, as there are numerous contributing factors from surface chemistry, mechanical properties, and sterics which are inherently compounded within the course of this study. 35000 30000 25000 20000 15000 10000 5000 0

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Figure 6. Cell adhesion assay for the PCL scaffolds, assessed at 4 hours following the initial seeding event. Note that ***p < 0.001, ****p < 0.0001, and n = 6.



25 The metabolic activity assay demonstrated that all of the polymeric scaffolds tested induced a slight reduction in metabolism, at a time point 24 hours following the initial seeding event, when compared against a tissue-culture treated control surface (Figure 7). Although this decrease in metabolism was determined to be statistically significant for the majority of polymeric scaffolds, the overall decrease observed for each fabrication method was relatively small in magnitude. The average reduction in metabolism was measured to be – 5.14 %, – 8.67 %, and – 5.99 % for ES, RJS, and AB products, respectively. The cause of slowed metabolism for hFOB cells may be attributed to a reduced rate of initial adherence within the three-dimensional structures.43 Corresponding cells that were seeded along the planar, tissue-culture treated control adhered at a faster rate, potentially leading to an earlier onset of a more robust cellular metabolism. 100000

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Figure 7. Metabolic activity assay for the PCL scaffolds, assessed at 24 hours following the initial seeding event. Note that *p < 0.05, **p < 0.01, ***p < 0.001, and n = 6. The proliferation assay demonstrated that effects stemming from the initial reduction in adherence rates and metabolic activity can still be observed within the polymeric scaffolds at 168 hours, but have been effectively surpassed by 336 hours (Figure 8). At 168 hours, the hFOB population within each of the polymeric scaffolds tested was lower than that measured for the control surface; with the average osteoblast population for only one scaffold, RJS 8 % PCL (w/v), falling within the lower standard deviation for said control. Conversely, at 336 hours, the hFOB populations within all of the polymeric scaffolds tested were not only above that measured for the control surface, but were beyond the upper


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26 standard deviation of said control. The osteoblast populations for all of the PCL scaffolds were over 50.0 % greater than that of the control at 336 hours, with the exception of the RJS 17 % PCL (w/v) scaffold which only bore a 33.6 % population increase in comparison to said control. While the heightened hydrophobic character may have inhibited the hFOB cells upon initial attachment to, and migration through, the PCL scaffolds, the increased surface area for sustained growth which is afforded by these three dimensional structures may explain the higher proliferation which is observed at longer time points. Phenotypic changes, stemming from altered cellular geometry (with adhesion complexes anchoring cells within a three-dimensional space, as opposed to along a planar surface) or substrate mechanosignaling, may have also resulted in heightened proliferation following the immediate inhibitions to adhesion and migration upon seeding (Figure 9).29;44,45 900000 800000 700000 600000 500000 400000 300000 200000 100000 0

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Figure 8. Cell proliferation assay for the PCL scaffolds, assessed at 168 hours and 336 hours following the initial seeding event. Note that **p < 0.01, ***p < 0.001, ****p < 0.0001, and n = 6.



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27 (A.)

(B.)

(C.)

Figure 9. Demonstration of critically-point dried hFOB cells within the PCL scaffolds at 168 hours following the initial seeding event. (A.) SEM imaging of hFOB cells within an ES 12 % PCL (w/v) scaffold, at 2,000x


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28 magnification. Observe that the cells have wrapped around individual microfibers, and have contacts extending through multiple planes in the scaffold. (B.) SEM imaging of hFOB cells within a RJS 12 % PCL (w/v) scaffold, at 2,000x magnification. Observe that the cells have spread over a large surface area, and that indentations from the microfibrous scaffolding are apparent along the cellular sheet. (C.) SEM imaging of hFOB cells within an AB 12 % PCL (w/v) scaffold, at 2,000x magnification. Observe that the cells have colonized the intraporous space of the nanofibrous scaffolding, through interstrand adhesions. Scale bars are 20 µm in length. Osteoblast Phenotype Evaluation – Following the evaluation that PCL scaffolds produced through ES, RJS, and AB were supportive for hFOB cell viability, the focus of experimentation shifted to determining osteoblast functionality within these structures. The bone formation phenotype was assessed among seeded hFOB cells by means of gene expression, alkaline phosphatase activity, and calcium deposition studies.

Gene expression studies measured relative mRNA levels for COL1A1 and COL1A2, which are translated in the cytoplasmic compartment to produce collagen, type one, alpha one and alpha two, respectively (Figure 10, A-B). These proteins represent key structural components of the dense, endogenous extracellular matrix (referred to as the osteoid) which is synthesized by fully-differentiated osteoblasts over the course of the bone formation process.46-48 Expression of COL1A1 and COL1A2 was up-regulated within all of the polymeric scaffolds tested relative to a tissue-culture treated control surface, with the exception of AB 8 % PCL (w/v) and AB 12 % PCL (w/v) which experienced effective downregulation. Gene expression studies measured the relative mRNA level for ALPL, which is translated in the cytoplasmic compartment to produce alkaline phosphatase, tissue non-specific isozyme (Figure 10, C). This enzymatic species is directly associated with the biomineralization of collagenous scaffolds during the



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29 course of bone formation, presumably by catalyzing the reduction of localized pyrophosphate and promoting higher inorganic phosphate concentrations, although the exact mechanism of action is still an area of ongoing study.49 Alkaline phosphatase is a membrane-bound protein, which is observed along the plasma membrane of fully-differentiated osteoblasts and matrix vesicles found within osteogenic sites.50,51 Expression of ALPL was up-regulated within all of the polymeric scaffolds tested relative to the control surface, with the exception of RJS 12 % PCL (w/v) and AB 8 % PCL (w/v) which experienced effective down-regulation. It is noted that ALPL expression for the AB 12 % PCL (w/v) scaffold was relatively low, but was still up-regulated respective to cells cultured along the control. Gene expression studies measured the relative mRNA level for SPP1, which is translated in the cytoplasmic compartment to produce osteopontin (Figure 10, D). This protein is excreted into the extracellular space, and serves as a regulatory species for the biomineralization of collagenous scaffolds during bone formation. Osteopontin acts to negatively control the growth of hydroxyapatite crystals along collagen fibrils by adhering to their exposed faces, thus sterically inhibiting further mineral deposition at these sites.52,53 Osteopontin adherence is driven through a combination of van der Waals forces and electrostatic attraction, the latter of which is ascribed to a series of extended, negatively charged, acidicresidue motifs along the length of the protein sequence.54 Expression of SPP1 was up-regulated within all of the polymeric scaffolds tested relative to the control surface, with the exception of AB 8 % PCL (w/v) which was analogous to the expression level of the tissue-culture treated control. It is noted that among the polymeric scaffolds tested in this experimentation, SPP1 expression was lowest for the AB products. Gene expression studies measured the relative mRNA level for BGLAP, which is translated in the cytoplasmic compartment to produce osteocalcin (Figure 10, E). This protein is excreted into the extracellular space, and serves as a hormone for paracrine and endocrine signaling.55-57 Osteocalcin is frequently viewed as a biomarker for active bone formation, as it is specifically produced and secreted by fully-differentiated osteoblasts.58 Expression of BGLAP was up-regulated within all of the polymeric


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30 scaffolds tested relative to the control surface, with the exception of AB 12 % PCL (w/v) which experienced effective down-regulation. BGLAP expression for the AB 8 % PCL (w/v) scaffold was also essentially analogous to that of the control surface. (A.)

Relative Expression

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(B.)

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31 (C.)

ALPL Relative Expression

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(D.)

Relative Expression

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(E.)

BGLAP 4 Relative Expression

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8% AB


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32 Figure 10. Gene expression studies for the PCL scaffolds, assessed at 336 hours following the initial seeding event. Relative gene expression is demonstrated for: (A.) COL1A1, (B.) COL1A2, (C.) ALPL, (D.) SPP1, and (E.) BGLAP. Relative gene expression was calculated by the ΔΔCT method, with GAPDH serving as the reference housekeeping gene. Note that n = 3. Upon review of the relative mRNA levels for COL1A1, COL1A2, ALPL, SPP1, and BGLAP, it would appear that there is a general trend indicating reduced osteogenic performance among the AB products, which represented the nanoscale fibrillar scaffolds tested within the scope of this study. However; mRNA level does not always directly correlate to an implied phenotypic response, as the bioactivity of translated proteins is largely modulated through a series of complex, post-translational modifications.59-61 These modifications may alter the inherent functionality of enzymatic species, or can alter the half-lives of proteins in general.62 In order to accommodate for this, alkaline phosphatase activity and calcium deposition were implemented as more direct assessments of osteogenic activity at the proteomic level. Alkaline phosphatase activity was measured at two time points, 168 hours and 336 hours. At 168 hours, half of the polymeric scaffolds assessed demonstrated enzymatic activity above the upper standard deviation of the tissue-culture treated control surface (Figure 11). The remaining polymeric scaffolds fell within the bounds of the standard deviation for said control, with only ES 12 % PCL (w/v), ES 17 % PCL (w/v), and RJS 8 % PCL (w/v) bearing slightly reduced alkaline phosphatase activity. Between 168 hours and 336 hours, there was a pronounced fluctuation in observed enzymatic activity. Only the ES 12 % PCL (w/v) and RJS 8 % PCL (w/v) scaffolds demonstrated an appreciable relative increase, with enzymatic activities 38.5 % and 51.4 % above the control surface at this later time point. Of the polymeric scaffolds that experienced a relative decrease in enzymatic activity, only ES 8 % PCL (w/v) and ES 17 % PCL (w/v) fell below the control at 336 hours. This reduction is specifically noteworthy for ES 17 % PCL (w/v), which already demonstrated considerably low enzymatic activity at 168 hours. It is also interesting to note that the enzymatic activity for both AB products was significantly above that of the control surface at both



33 times points, which does not correlate with the down-regulated ALPL expression observed within these scaffolds. 3 2.5 2 1.5 1 0.5 0

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Figure 11. Alkaline phosphatase bioactivity within the PCL scaffolds, assessed at 168 hours and 336 hours following the initial seeding event. Note that *p < 0.05, **p < 0.01, ***p < 0.001, ****p < 0.0001, and n = 6. Calcium deposition was similarly measured at two time points, 168 hours and 336 hours, and demonstrated the most staggering results observed within the course of this study (Figure 12). At 168 hours, all polymeric scaffolds tested had an over 3,000 % increase in deposited calcium relative to the tissue-culture treated control surface. This heightened biomineralization occurred in spite of potential complications stemming from the initial hindrances to adherence, metabolism, and proliferation. Heightened biomineralization was observed among the ES 12 % PCL (w/v), ES 17 % PCL (w/v), and RJS 8 % PCL (w/v) scaffolds, which were shown to have low alkaline phosphatase bioactivity at 168 hours; as well as both AB products, which demonstrated a general trend of down-regulating the expression of gene sequences related to osteogenesis. At 336 hours, all polymeric scaffolds tested had an over 5,000 % increase in deposited calcium relative to the tissue-culture treated control surface. This included the ES 8 % PCL (w/v) and ES 17 % PCL (w/v) scaffolds, which were previously shown to experience a reduction in relative alkaline phosphatase bioactivity between the 168-hour and 336-hour time points.


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34

Figure 12. Calcium deposition within the PCL scaffolds, assessed at 168 hours and 336 hours following the initial seeding event. Note that ****p < 0.0001, and n = 6. Observations from alkaline phosphatase activity and calcium deposition studies demonstrate that a systematic review of mRNA transcriptional levels may not be sufficient for evaluating the osteogenic potential of engineered biomaterials, such as the PCL scaffolds tested herein. Perhaps most interesting is that calcium deposition, which most directly measures biomineralization, did not maintain trends observed during the assessment of alkaline phosphatase bioactivity. This seemingly indicates that utilizing alkaline phosphatase activity as the sole protein-based assay for bone formation may be similarly insufficient, as numerous additional biomolecular species (including osteopontin and osteocalcin, as well as osteonectin, osteomodulin, and bone sialoprotein) will influence the progression of orthopedic biomineralization.63-66 Bacterial Viability Evaluation – The colony forming unit assay was utilized to determine the susceptibility of the polymeric scaffolds to infection from gram-positive (S. aureus) and gram-negative (P. aeruginosa) bacterial strains. Results indicated that the nanofibrillar AB products experienced reduced bacterial colonization when inoculated with either S. aureus or P. aeruginosa, in comparison to microfibrillar scaffolds produced via ES or RJS (Figure 13). With regards to S. aureus, AB products had a reduction of – 78.3 % and – 31.1 % colony forming units when compared with ES products at 8 % PCL (w/v) and 12 % PCL (w/v), respectively;



35 and had a reduction of – 71.1 % and – 68.1 % colony forming units when compared with RJS products at 8 % PCL (w/v) and 12 % PCL (w/v), respectively. With regards to P. aeruginosa, AB products had a reduction of – 90.5 % and – 79.6 % colony forming units when compared with ES products at 8 % PCL (w/v) and 12 % PCL (w/v), respectively; and had a reduction of – 96.1 % and – 94.2 % colony forming units when compared with RJS products at 8 % PCL (w/v) and 12 % PCL (w/v), respectively. The resistance of AB products to bacterial colonization (without utilizing small molecule antibiotics or any comparable biomolecular additives) may potentially be attributed to the nanoscale dimensions of its composing fibers, which provide relatively small adhesion sites for opportunistic bacterial species in comparison to the microfibrous ES and RJS products. Previous studies have demonstrated that bacterial adhesion to submicron fibers can result in altered cellular morphology, which may translate to phenotypic changes and ultimately reduced viability.67 While these adhesion sites may not promote prokaryotic expansion within the AB scaffolds, this does not appear to inhibit osteoblast proliferation and subsequent osteogenesis.6870

The larger characteristic length scale of osteoblasts appears to permit several focal contact points

between multiple local fibers, resulting in the interstrand binding and intraporous colonization observed within this study (Figure Two). (A.)

Bacteria Density (CFU/mL)

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50000000

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36 (B.) P. aeruginosa 600000000 Bacteria Density (CFU/mL)

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* * *

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Figure 13. Colony forming unit assay for (A.) gram-positive S. aureus and (B.) gram-negative P. aeruginosa bacterial strains within the PCL scaffolds, assessed at 24 hours following the initial inoculation event. Note that *p < 0.05, **p < 0.01, ***p < 0.001, and n = 9. Conclusions: Three fabrication methods were investigated for the production of readily translatable, orthopedic implants, based on PCL scaffolds: ES, RJS, and AB. While ES and RJS fabrication produced microfibrillar structures, AB produced nanofibrillar structures. Although all polymeric scaffolds synthesized in this study were tested without any further modifications, these structures can be readily modified through plasma treatment or through the facile incorporation of nanomaterials (such as carbon nanotubes and hydroxyapatite nanoparticles) upon synthesis. While all of the polymeric scaffolds tested within the course of this study demonstrated initial hindrances to osteoblast adherence, metabolism, and proliferation, all ES, RJS, and AB products still supported mammalian cell viability, as seeded hFOB cells were capable of out-proliferating an analogous control population (grown along a planar, tissue-culture treated surface) at long time points. While reviewing osteogenic potential within the seeded PCL scaffolds, initial gene expression studies demonstrated that nanofibrous AB products may be at a significant disadvantage for osteoid production and subsequent biomineralization, in comparison to the both the microfibrillar scaffolds and the planar control surface. This projected behavior was not observed upon the evaluation of alkaline phosphatase



Page 56 of 65

37 bioactivity and calcium deposition, as all ES, RJS, and AB products demonstrated significantly heightened biomineralization relative to the tissue-culture treated control. The discrepancy in results from the gene expression studies with those observed for alkaline phosphatase bioactivity and calcium deposition assays was attributed to that fact that relative mRNA levels do not always translate to direct phenotypic changes. A complex network of post-translational modifications may alter the activity and longevity of enzymes and protein species within a cell; and as these modification pathways have not been fully elucidated for the osteoid synthesis and biomineralization processes in general, they were not directly assessed within the scope of the present study. Although all of the polymeric scaffolds tested within the course of this study supported osteoblast viability, and promoted the bone formation phenotype, the nanofibrous AB products demonstrated significantly reduced bacterial colonization, for both gram-positive and gram-negative bacterial strains, when compared to the microfibrillar ES and RJS products. This effect was achieved without the incorporation of small molecule antibiotics, antimicrobial peptides, or reactive nanomaterials within the scaffold design. The reduced bacterial colonization observed for AB products was attributed to the nanoscale dimensions of its comprising fibers, as previous studies have demonstrated bacteria adopting pronounced morphological changes when adhering to sub-micron diameter fibers – with these morphological changes ultimately resulting in an altered phenotype and observed viability. While unmodified PCL scaffolds have demonstrated potential as osteoconductive implants for musculoskeletal defects, with a simplistic design promoting both industrial scale-up and regulatory approval, future studies will be directed at determining how post-manufacture processing and nanomaterial incorporation can further tailor implant properties for specific tissue-regeneration challenges. Plasma treatment provides a mechanism to tailor the hydrophilicity of PCL scaffolds, and thus promote osteoblast infiltration and viability upon implantation at a wound site. This treatment also provides a pathway for the functionalization of polymeric scaffolds with a diverse range of


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38 biomacromolecular species, such as bone morphogenic proteins, adhesion peptides, and antimicrobial peptides. The introduction of nanomaterials within PCL scaffolds similarly presents a means to further modulate their mechanical and degradation properties, thus influencing mechanosignaling to adherent cells and altering the substrate’s post-implantation persistence. Both plasma treatment and nanomaterial incorporation will be investigated in future studies with PCL-based orthopedic scaffolds. Future studies with these PCL scaffolds will also focus on the mechanism of promoted osteogenesis, quantifying the presence and bioactivity of alkaline phosphatase, osteopontin, osteocalcin, osteonectin, osteomodulin, and bone sialoprotein among other protein species. Protein expression levels will once again be projected by qPCR analysis, but with a direct assessment of abundance and posttranslational modification completed via quantitative proteomic techniques, such as two-dimensional gel electrophoresis and mass spectroscopy. An analogous assessment will be applied bacterial species seeded within PCL scaffolds, in order to assess select biomarkers for prokaryotic metabolism and viability, and compare these against observed morphological changes upon adherence to nanofibrillar and microfibrillar scaffolds. As the potential clinical advantages of such simplistic orthopedic implants are readily apparent, capable of promoting biomineralization with reduced bacterial infection, future research on these PCL scaffolds will focus on the AB fabrication method. A greater range of gram-positive and gram-negative bacterial strains will be tested, including: Staphylococcus epidermidis, Streptococcus pneumoniae, methicillin-resistant Staphylococcus aureus, and MDR Escherichia coli. At present, the AB scaffolds demonstrate the greatest potential for applications treating musculoskeletal defects, by promoting the regeneration of native bone tissue with reduced propensity for bacterial infection.



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39 Acknowledgements: Many thanks to William Fowle and Scott McNamara for training on the S – 4800 SEM (Hitachi) and Phoenix 150 Contact Angle Analyzer (Surface Electro Optics), respectively.

Institution Identification: Northeastern University, Boston, Massachusetts, 02115, USA. Brigham and Women’s Hospital, Harvard Medical School, Boston, Massachusetts, 02115, USA. Harvard – MIT Health Sciences and Technology, Massachusetts Institute of Technology, Cambridge, Massachusetts, 02139, USA. Universidade do Vale do Paraíba, São José dos Campos, São Paulo, 12244 – 000, Brazil. Universidade Federal do Piauí, Teresina, Piauí, 64049 – 550, Brazil.

Funding Sources: This work was supported by funding from Northeastern University.


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Micro-Nanofibrillar Polycaprolactone Scaffolds as Translatable

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