Influence of Cross-Linking on the Physical Properties and Cytotoxicity

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Influence of cross-linking on the physical properties and cytotoxicity of polyhydroxyalkanoate (PHA) scaffolds for tissue engineering Alex C Levine, Angelina Sparano, Frederick F Twigg, Keiji Numata, and Christopher T. Nomura ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.5b00052 • Publication Date (Web): 27 May 2015 Downloaded from http://pubs.acs.org on May 29, 2015

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ACS Biomaterials Science & Engineering

Influence of cross-linking on the physical properties and cytotoxicity of polyhydroxyalkanoate (PHA) scaffolds for tissue engineering Alex C. Levine1, Angelina Sparano1, Frederick F. Twigg1, Keiji Numata2, Christopher T. Nomura1,3* 1

Department of Chemistry, State University of New York − College of Environmental Science and Forestry, 1 Forestry Drive, Syracuse, New York 13210, United States 2 Enzyme Research Team, RIKEN, 2-1 Hirosawa, Wako, Saitama 351-0198, Japan 3 Center for Applied Microbiology, State University of New York − College of Environmental Science and Forestry, 1 Forestry Drive, Syracuse, New York 13210, United States

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*[email protected]

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Telephone: 315-470-6854; Fax: 315-470-6856

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Keywords: polyhydroxyalkanoates, biomaterials, Escherichia coli, cross-linking, thiol-ene click

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chemistry, human mesenchymal stem cells, tissue engineering, scaffold

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Abstract

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In this study an unsaturated copolyester, poly[(R)-3-hydroxybutyrate-co-(R)-3-hydroxy-10-

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undecenoate] (PHBU) was produced by an engineered strain of Escherichia coli, cross-linked

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via thiol-ene click chemistry, and analyzed for improved physical properties and biocompatibility

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with human mesenchymal stem cells.

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tensile strength of over 200% to 26.2 MPa was observed, resulting in a material with physical

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properties closer to those relevant for soft tissue replacement.

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chemically cross-linked polyester did not exhibit significant cytotoxicity toward human cells after

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chemical modification. The chemically modifiable copolyester described here could potentially

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be used as a replacement for an assortment of tissues currently without viable material

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alternatives in the field of tissue-engineering.

By cross-linking the PHBU polymer, an increase in

Results showed that this

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Introduction

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Biomedical scaffolds are materials that are used to replace or regenerate living tissue1

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by providing support for cell attachment and growth. Due to the great variety of cell types and

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complex nature of cell attachment2, an ideal scaffold material should be tailorable to encourage

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the production of or mimic the extracellular matrix of target tissues3 in order to facilitate

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compatibility and growth, and in the case of stem cells, differentiation4. The main considerations

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when selecting the appropriate material for a tissue-engineering scaffold are architecture,

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tissue/cyto-compatibility, bioactivity, and mechanical properties1,5. The architecture of the

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scaffold should be highly porous to allow transport of nutrients and vascularization, although not

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so porous as to compromise the strength of the material6,7,6. The scaffold should also

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biodegrade in vivo while simultaneously being replaced by cells of the target tissue8. In order to

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support cell attachment, growth, and proliferation, the scaffold material must be compatible with

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cell types and components both in vitro and in vivo9. In vitro compatibility is relevant when

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seeding the scaffolds with cells prior to implantation and in vivo compatibility is crucial to prevent

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immune response and rejection of the implant. The bioactivity of the scaffolds must also be

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considered, as cell attachment and morphology can be effected by the scaffold topology2. In

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addition to surface chemistry, the smoothness or roughness of the scaffold surface can greatly

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affect cellular attachment. If the scaffold is serving as a delivery method for biological signals to

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stimulate growth10, then the release of the compound must be stringently controlled for

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therapeutic efficacy and non-toxicity. Finally, the mechanical properties of the scaffold must

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match that of the target tissue3. Scaffolds with mismatching mechanical stiffness or elasticity will

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have decreased performance in situ and the implant itself may be prematurely degraded or

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irreparably damaged11,12.

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Polymers used for tissue-engineered scaffolds vary greatly, and include both synthetic

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and natural polymers such as proteins, polysaccharides, and polyesters13,14. Biological polymers

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are of great interest as tissue engineering materials as they have greater biocompatibility and

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biodegradability compared to synthetic polymers, though the mechanical properties of these

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biopolymers are often inferior to synthetic materials. Collagen is a well-studied biopolymer in

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tissue-engineering due to its bioactivity and inherent biocompatibility15,16. However, collagen

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alone cannot always meet the mechanical requirements for tissue replacement, and it is

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common for collagen to be blended with other polymers to create composite materials.17 In

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these composite materials, the role of collagen is to increase bioactivity and cellular attachment,

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whereas the other purpose of the other polymer is to provide suitable strength to the material to

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prevent damage and improve application. Collagen composites with polylactic acid, a type of

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biopolyester, have been used for vascular grafts where the collagen provides a surface that

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cells can readily attach to, and the biopolyester serves to maintain the shape of the tissue.18

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Polysaccharides such as cellulose are also of interest, and it has recently been shown that the

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native cell-wall structure in apples can be used as a tissue scaffold19. In the realm of polyesters,

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polylactic acid (PLA)20, polycaprolactone (PCL)13,21, and polyhydroxyalkanoates (PHA)22,23,24

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have been well studied. PHAs are of particular interest as they have a greater range of physical

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properties as compared to other biopolyesters, with applications such as heart valves already

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showing promising results in animal models25.

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Previous studies have demonstrated the effect of repeating unit composition on the

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material properties of PHA polymers and have shown that PHA polymers comprised of short-

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chain-length (SCL) repeating units of 3-5 carbons behave as thermoplastics and PHA polymers

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comprised of medium-chain-length (MCL) repeating units of 6-14 carbons behave as

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elastomers. Copolymers of SCL and MCL repeating units have physical properties ranging from

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SCL to MCL homopolymers. In addition, by controlling the ratio of SCL to MCL repeating units,

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the material properties can cover a wide range of physical properties26,27,28,29. Recently our lab

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engineered a strain Escherichia coli LSBJ (E. coli LSBJ) to produce PHA homopolymers and

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copolymers with defined repeating compositions. Strictly defining the repeating unit composition 4


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of PHA polymers and copolymers allows for stringent control over PHA physical properties such

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as mechanical strength, thermal properties, and hydrophobicity28,29. These PHA materials may

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be suited for tissue-engineering scaffolds based on the required stiffness for the application and

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biodegradability. Additionally, for the first time, PHA polymers produced from this strain of

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bacteria can be programmed to contain specific quantities of olefin functional groups, which can

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be subsequently modified via synthetic chemistry to provide further diversification of material

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properties. While other works have shown the potential to chemically modify PHAs

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previous studies were not focused on tissue-engineering applications and used PHA polymers

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where repeating unit compositions were highly variable due to the native metabolism of the

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bacterial strain used to produce said polymers.

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, these

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The goal of this work was to produce a strong and flexible PHA scaffold of PHA with

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inherent biocompatibility.30,32,33 In this work, we have produced and characterized a new PHA

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copolymer for tissue engineering scaffolds, poly[(R)-3-hydroxybutyrate-co-(R)-3-hydroxy-10-

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undecenoate] (PHBU). PHBU is a short-chain-length/medium-chain-length (SCL/MCL) PHA

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copolymer that contains a terminal alkene functional group in the side chain of its MCL

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repeating unit. This degree of unsaturation was used as a site to strategically improve the

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physical properties of the PHBU scaffolds using thiol-ene click chemistry. The benefits to the

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physical properties of polymers via cross-linking have been well established34,35, with observed

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improvements to tensile strength, stiffness, and cell attachment. PHA cross-linked by thiol-ene

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click chemistry has tailorable physical characteristics dependent on the density of cross-linking,

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and can be controlled by varying the concentration of the cross-linker.

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here demonstrates that thiol-ene click chemistry can be employed to chemically cross-link PHA,

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resulting in a PHA scaffold with significantly enhanced tensile strength and insignificant effects

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on cytotoxicity.

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engineered scaffolds that require stiffer materials, such as cartilage and ligament

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replacements34,35,36,37.

The work presented

It is anticipated that this cross-linked PHA scaffold will find use in tissue-

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Experimental

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E. coli LSBJ. The SCL/MCL copolymer PHBU was produced as described for poly[(R)-3-

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hydroxybutyrate-co-(R)-3-hydroxyoctanoate] by Tappel et. al29. The strain used for polymer

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production was E. coli LSBJ, a derivative of E. coli LS5218 with deletion of genes fadB and

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fadJ, and harboring a plasmid containing the genes for an (R)-specific enoyl-CoA hydratase

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(phaJ4) and PHA synthase [phaC1(STQK)]29.

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Erlenmeyer flask containing 50 mL Lennox broth (LB) and 50 mg/L kanamycin, which was

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inoculated with E. coli LSBJ and incubated for 16-18 h at 30°C in an orbital shaker set at 250

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rpm. For polymer production, 1 mL of this culture was used to inoculate 500 mL baffled flasks

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containing 100 mL LB medium, 50 mg/L kanamycin, 4 g/L Brij-35 (surfactant), and 2 g/L of fatty

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acids.

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Following, cells were harvested with centrifugation for 15 minutes at 3716 x g and 22 °C, and

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the supernatant was discarded. Cell pellets were washed with 70% ethanol to remove residual

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fatty acids, collected again by centrifugation under the same conditions, and washed with

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nanopure water. After washing with water, the cells were collected by centrifugation once more

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and resuspended in 15 ml of water before freeze drying. After drying by lyophilization, the cells

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were suspended in methanol (22 mL/g dried cell mass) and gently stirred for 5 minutes at 22°C.

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Afterwards, cells were collected by centrifugation and washed with nanopure water. Cells were

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again dried by lyophilization, and PHAs were purified via Soxhlet extraction followed by

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methanol precipitation. Soxhlet extractions were performed in 120 mL of chloroform for 5 hours,

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and then the solution was transferred to a glass petri dish where the solvent was evaporated

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under ambient conditions for >6 hours. The resulting film was dissolved in a minimal amount of

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chloroform, and precipitated into a 10-fold larger volume of methanol. The final white powder

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was stored at 22°C in the dark until scaffold fabrication.

Poly[(R)-3-hydroxybutyrate-co-(R)-3hydroxy-10-undecenoate] (PHBU) Production using

Starter cultures consisted of a 250 mL

Cultures were grown at 30° C for 48 hours in an orbital shaker set at 250 rpm.

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Molecular weight and repeating unit content determination

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Molecular weight was determined using a Shimadzu GPC system. Samples were

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prepared for GPC by transferring 1-2 mg of polymer to a 5 mL glass vial and dissolving the

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material in chloroform to a final concentration of 0.7 mg/mL by heating at 50 °C.

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dissolved, 1 mL of solution was filtered through a 0.45 µm PTFE syringe filter into a 2 mL GPC

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vial. Sample volumes of 50 µL were injected into a LC-20AD liquid chromatograph equipped

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with a SIL-20A autosampler, CTO-20A column oven, and an RID-10A refractive index detector

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(Shimadzu). Chromatography was performed using a 8 × 50 mm styrenedivinylbenzene (SDV)

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guard column (5 µm particles; Polymer Standards Service) and a 8 × 300 mm SDV analytical

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column (5 µm particles; mixed bed porosity; max molecular weight 1E6 Da; Polymer Standards

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Service product sda083005lim). The mobile phase consisted of chloroform at a flow rate of 1

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mL/min and the temperature was maintained at 40 °C. Analysis was performed using Shimadzu

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LCsolution software. The repeating unit content of PHBU was determined by

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spectroscopy using a Bruker AVANCE 600 spectrometer. Samples consisting of 15 mg of

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polymer were dissolved in 1 mL of deuterated chloroform. Spectra were processed using

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TOSPIN v1.3 from Bruker BioSpin. The ratio of the alkene proton signal at 5.8 ppm to the

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polymer backbone stereocenter at 5.3 ppm was taken as the MCL portion of the copolymer. A

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series of PHBU copolymers with varying MCL content were generated to produce a standard

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curve for precise PHBU synthesis. From this data, a PHBU copolymer with 8% MCL was

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produced for scaffold fabrication.

Once

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HNMR

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PHA scaffold fabrication. PHA scaffolds were fabricated using a combined method of salt-

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leaching20,21,40 and thermally induced phase separation1,10,40. To a 3 mL glass vial, 0.10 g of

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PHBU (8.4 x 10-5 moles of alkene) and 1 mL of 1,4-dioxane were added, and the polymer was

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dissolved by stirring and heating at 60 °C for 30 minutes. To this solution, 1.0 g of sodium

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chloride was added, and the resulting heterogeneous mixture was transferred to a custom-made

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mold consisting of two microscope slides separated by a 1 mm thick teflon spacer. The 7



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dimensions of the mold's chamber were 60 mm, 25 mm, and 1 mm (L x W x T). The samples

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were frozen at -80°C for 1 hour, and then the solvent removed by lyophilization for 24 hours.

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The resulting PHBU/salt composite was placed in 300 mL Milli-Q water with stirring at 25 °C for

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4 hours, with complete water changes after each hour. The composite was soaked for an

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additional 48 hours, with water changes every 24 hours. The resulting PHBU scaffold was dried

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under vacuum for 48 hours.

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PHBU scaffold cross-linking. PHBU contains a terminal alkene functional group as part of

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the polymer side-chain. This site can be chemically modified via thiol-ene click chemistry41.

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Thiol-ene click chemistry is a radical mediated coupling reaction between alkene and thiol

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functional groups, resulting in a covalent thioether linkage in a fast, efficient, and water/oxygen

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insensitive reaction. The PHBU copolymer was cross-linked using pentaerythritol tetrakis (3-

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mercaptopropionate) (PETMP) and radical initiator 2,2-dimethoxyphenylacetophenone (DMPA),

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both purchased from Sigma-Aldrich. For fully cross-linked samples, 8.0 µL of PETMP (2.1 x 10-

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5

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The amount of PETMP was chosen to be 25% the concentration of available alkene functional

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groups in PHBU due to PETMP’s four thiol functional groups. This rationale was based on the

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criteria of maximizing cross-link density and reducing the amount of unreacted PETMP in the

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PHBU scaffolds. Samples were also made using half PETMP concentration at 4.0 µL. After

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transferring the polymer solution, the mold was placed under a UVP UVGL-58 handheld lamp

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and irradiated with 365 nm UV light for 1 hour. Following irradiation, the samples were frozen

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and treated identically to non-cross-linked samples with treatment by lyophilization, salt

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leaching, and drying.

moles) and 1 mg of DMPA were added to the PHBU solution prior to transfer into the mold.

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Assessment of cross-linking by Fourier transform infrared spectroscopy (FT-IR). FT-IR

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was utilized to determine the degree of cross-linking in the PHBU polymer by measuring the

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change in stretching vibration at 1641cm-1, which corresponded to the alkene functional group.

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The reduction of this signal was taken as evidence for formation of new thioether linkages via 8


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the thiol-ene click reaction. Data was collected using a Bruker Tensor 27 spectrometer equipped

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with an attenuated total reflection (ATR) stage. Each sample received 32 scans and data was

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extracted in transmittance mode with OPUS 6.5 software. Data was re-plotted using Microsoft

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Excel 2007.

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Gel soluble fraction. To determine the extent of the cross-linked network within the scaffolds,

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the gel fraction35 of each sample was measured after incubation in chloroform. PHBU samples

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of 5-10 mg were weighed, and each placed into 3 mL glass vial containing 2 mL of chloroform.

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After incubation at 25 °C for 48 hours the solvent was removed, and the swollen PHA polymer

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was rinsed with a fresh 1 mL portion of chloroform. At this time, the cross-linked portion of the

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PHA scaffold was dried under vacuum for 48 hours. The gel fraction was determined by:

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Gf = W f / W 0 x 100

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Where Gf is the weight percentage of cross-linked PHBU, W f is the weight of the sample after

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chloroform extraction, and W 0 is the initial weight of the sample.

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Cell culture and viability. To assess compatibility with living cells, human mesenchymal

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stem cells (hMSC)4, a type of multipotent stromal cell able to differentiate into a variety of cell

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types, were used to seed the PHBU scaffolds. In order to determine whether cross-linking had

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an effect on cytotoxicity, cultures of hMSC seeded onto the PHBU scaffolds were examined.

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The scaffolds were cut into 5x5x1 mm squares, and attached to the bottom of the wells in a

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polystyrene 96 well plate using a small amount of chloroform, and dried under vacuum for 48

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hours. For sterilization, scaffolds were soaked in 70% ethanol for 2 hours, after which the

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ethanol solution was removed and the scaffolds were washed twice with 200 µL portions of D-

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PBS(-) buffer (Wako, cat. no. 045-29795). Finally, the scaffolds were incubated in 200 µL PBS

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buffer for 18 hours in a sterile hood under sterilizing UV light. hMSC were cultured in 75 cm2

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polystyrene flasks using hMSC Expansion Media (StemXVivo™), designed specifically for

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human mesenchymal stem cells. Cultures contained 7-10 mL of hMSC Expansion Media and

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where incubated at 37°C with 5% CO2 for 48 hours. hMSC Expansion Media was replaced with 9



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a fresh portion after 24 hours, at this time also containing the antibiotic-antimycotic Anti-Anti™

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(Life Technologies, cat. no. 15240-062). Once the cultures reached 80-90% confluence, the

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cells were removed from the flasks by trypsin digestion and counted using a hemocytometer. To

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each well containing a scaffold, 8,000 cells in 200 µL of hMSC expansion media were added,

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and incubated at 37°C with 5% CO2 for 24 hours to allow for cell attachment. After 24 hours,

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the medium was exchanged with a fresh 200 µL portion of hMSC expansion media containing

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antibiotic-antimycotic, and the cells were allowed to grow for an additional 24 hours. Viability

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was measured by (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-

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2H-tetrazolium) (MTS) Assay (Promega), where at 48 hours after hMSC seeding, 20 µL of the

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MTS reagent was added to each well containing a scaffold, and incubated at 37°C with 5% CO2

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for 1 hour. The supernatant was collected, and its absorbance measured at 490 nm using a

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Molecular Devices SpectraMax M3 plate reader.

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Viability was also observed by Invitrogen's Live/Dead™ assay (Life Technologies). After the 48

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hour incubation, the medium was removed from the scaffolds, and 100 µL of a solution

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containing calcein AM (2 mM) and ethidium homodimer-1 (4 mM) was added. After 30 minutes

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at 37°C with 5% CO2, the scaffolds were removed from the 96-well plate, placed on a glass

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microscope slide, and viewed using a ZeissLSM 700 confocal microscope. Excitation

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wavelengths of 488/555 nm were used to observe the fluorescence of live (green) and dead

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(red) cells.

Each sample consisted of 3-5 replicates.

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Scanning electron microscopy. PHBU scaffolds were analyzed using a JCM-6000

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NeoScope Benchtop scanning electron microscope set to backscattered electron image (BEI) at

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10 kV accelerating voltage, high vacuum, and high probe current. To observe hMSC adhered to

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the scaffolds, samples from the Live/Dead® assay were fixed with a 3% glutaraldehyde solution

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for 4 hours, followed by dehydration using a series of aqueous ethanol solutions of 70%, 80%,

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90%, 95%, and 99.5%. Each round of ethanol dehydration was performed for 1 h at 25 °C, at

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which point the solution was completely replaced by the next higher concentration in the series. 10


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Following the final dehydration step, scaffolds were dried under vacuum for 24 hours.

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Preparation for microscopy involved mounting samples on carbon tape and sputter coating for

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30 seconds with gold.

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Analysis of mechanical properties. Tensile strength was measured using a Shimadzu EZ-

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LX HS universal tester equipped with a 500 N load cell.

PHBU samples with dimensions

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averaging 20x2x0.1 mm were pulled at a constant rate of 2 mm/min until break. The tensile

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strength, Young's modulus, and elongation to break were determined based on the generated

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stress/strain curves.

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Results

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with unsaturated repeating unit content that could be cross-linked to improve the mechanical

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properties of the material to match those of soft tissues such as cartilage. This material requires

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a high amount of stiffness and enough flexibility to prevent brittleness. Based on a previous

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study28, a PHBU copolymer with 8% MCL content and 92% SCL content was produced for its

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balanced physical properties and its ample amount of MCL content for cross-linking. The fatty

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acids chosen for this polymer were butyric acid (cultured as sodium butyrate to reduce toxicity)

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and 10-undecenoic acid, which resulted in the copolymer poly[(R)-3-hydroxybutyrate-co-(R)-3-

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hydroxy-10-undecenoate] (PHBU). A series of PHBU copolymers were produced with a range

284

of MCL repeating unit contents in order to determine the parameters to produce copolymers

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with a final desired ratio of repeating units. A standard curve depicting the ratio of fatty acid

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concentration to repeating unit composition for PHBU was generated (Figure 1a).

287

results allowed for the design of a feeding strategy to produce PHA with 8% MCL content by

288

adjusting the fatty acid ratio in the culture medium to the correct specifications based on the

289

established trend.

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spectroscopy (Figure 1b), where the ratio of the alkene proton signal at 5.8 ppm to the polymer

Production and characterization of PHBU copolymer. A PHA copolymer was designed

The repeating unit content of PHBU was determined by

These

1

HNMR

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backbone stereocenter at 5.3 ppm was taken as the MCL portion of the copolymer. The yield of

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PHBU was approximately 50 wt% of the cell dry weight, which was similar to previous

293

studies.28,29 The number average molecular weight (Mn) of PHBU was estimated to be 74.3

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kDa by gel permeation chromatography (Table 1), and this number was used in the

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stoichiometric calculations for the thiol-ene click reaction.

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Fabrication of PHBU scaffolds. Porous PHBU scaffolds were produced with both small pore

297

sizes of 30-50 µm and large pore sizes of 200-400 µm (Figure 2). The smaller pores resulted

298

from the thermally induced phase separation of PHBU with 1,4-dioxane, and the larger pores

299

were derived from the salt-leaching of sodium chloride. It was anticipated that the smaller pore

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size would enhance cellular attachment and nutrient transport, while the larger pore size could

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promote vascularization of the scaffold if necessary. Analysis by scanning electron microscopy

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(SEM) revealed a uniform dispersion of pore sizes, with the walls of large pores containing the

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smaller pore size. This architecture should benefit transport of nutrients and waste materials, as

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the large pores were interconnected by a network of smaller pores present throughout the

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scaffold. These scaffolds were soft to the touch and sponge-like, with much greater flexibility

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than the PHBU films of the same repeating unit composition.

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Gel fraction of PHBU scaffolds. After scaffold formation, the extent of polymer participation

308

in the cross-linked network (gel fraction) was determined by swelling the material in chloroform

309

(Scheme 1). In scaffolds that contained a 1:4 ratio of PETMP to alkene functional groups, the

310

gel fraction was determined to be 98%. This ratio was chosen in order to maximize the amount

311

of cross-linking in the scaffold, due to the four available thiol groups in PETMP. Scaffolds were

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also prepared with half the amount of PETMP cross-linker, and this resulted in scaffolds with a

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gel fraction averaging 50%.

314

FT-IR analysis of cross-linked PHBU. To assess the degree of conversion to the cross-

315

linked product, PHBU and cross-linked PHBU materials were analyzed by FT-IR (Figure 3).

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This qualitative measure of alkene content in the final polymer showed that the signal 12


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corresponding to the alkene functional group in the side-chain of PHBU was reduced after the

318

cross-linking reaction. Samples containing an equimolar concentration of PETMP thiol groups to

319

PHBU alkenes (PHBUX98%) showed a reduction in the 1641 cm-1 signal by more than 90%,

320

whereas samples containing half that concentration of PETMP (PHBUX50%) had a reduction in

321

the alkene signal of approximately 50%. This result confirms that the concentration of cross-

322

linking can be controlled by the precise addition of the PETMP cross-linking agent.

323

Viability of human mesenchymal stem cells (hMSC) on PHBU scaffolds. After the

324

scaffolds were fabricated and sterilized, each one was seeded with hMSC and incubated for a

325

total of 48 hours. Cell viability measured by Live/Dead™ assay using confocal fluorescence

326

microscopy showed similar amounts of live and dead cells on the cross-linked scaffolds as

327

compared to non-cross-linked scaffolds (Figure 4).

328

showed no significant difference in metabolic activity between the positive control with no

329

scaffold and PHBU scaffold (p=0.82). No significant difference was also observed when

330

comparing the PHBU scaffold and cross-linked PHBU scaffold (p=0.54), (Figure 5). Analysis of

331

the cell-seeded PHBU scaffolds by SEM (Figure 6) revealed that greater amounts of

332

extracellular matrix formed on small pore size in the 30-50 µm range, indicating that this pore

333

size promoted higher amounts of hMSC attachment and growth.

Cell viability measured by MTS assay

334

Mechanical strength analysis of cross-linked PHBU. The process of cross-linking PHBU

335

resulted in significant increases to tensile strength. The increase of gel fraction in cross-linked

336

PHBU correlated to an increase in tensile strength. PHBU cross-linked to 50% gel fraction

337

exhibited an increase in tensile strength from 8.5 to 11.2 MPa, whereas PHBU cross-linked to

338

98% gel fraction showed an increase to 26.2 MPa. The strain to failure did not change

339

significantly between the three samples, though there was a decrease in Young's modulus with

340

increased cross-link density (Table 2). This decrease in modulus is likely the plasticizing effect

341

of the PETMP cross-linking agent. The ester functional groups of PETMP make it similar in

342

structure to known plasticizers for PHAs42, such as fatty acids and triglycerides. Notably, the 13



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343

shape of the PHBU stress/strain curve (Figure 7a) differs from the cross-linked PHBU

344

stress/strain curve (Figure 7b), where the cross-linked sample changed to closely resemble a J-

345

shape stress/strain curve.

346

Discussion

347

Tissue-engineering requires that the material used to fabricate the scaffold closely match

348

that of the target tissue. In addition to meeting physical requirements, the material must also be

349

biocompatible and biodegradable. PHAs are suited for tissue-engineering applications due to

350

their inherent biocompatibility, biodegradability, and diverse set of physical properties available

351

among SCL-PHA, MCL-PHA, and PHA copolymers. While these attributes are found in other

352

biopolymers such as PLA and PCL, PHAs have the advantage of chemical modification to

353

further tune the characteristics of the polymer to the target tissue. This secondary level of

354

control in physical properties allows for PHAs to be tuned for specific tissue replacements in a

355

way not possible with other biopolymers.

356

PHAs are typically limited to very soft tissue applications due to their low tensile

357

strength43. The cross-linked PHBU scaffolds produced in this study have substantial increases

358

in tensile strength compared to other SCL/MCL PHAs, with mechanical properties similar to

359

connective tissues such as fibrocartilage and ligaments (Table 3). In addition to changes in

360

physical properties, the shape of the stress/strain profile of cross-linked PHBU now more closely

361

resembles that of biological tissue. MCL PHAs have large deformations when a stress close to

362

their ultimate tensile strength is applied. During this period of deformation, dramatic increases in

363

elongation are seen with little or no increase in measured tensile strength. Contrary to this,

364

cross-linked materials have a J-shape stress-strain curve44,45, where during the period of high

365

elongation, there is a distinct increase in measured tensile strength. This type of profile was

366

observed in the cross-linked PHBU stress/strain curve of Figure 7b, where increased strain was

367

coupled with increased tensile strength over the entire curve. This result indicated that the

14


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368

cross-linked PHBU had deformation behavior similar to biopolymers like collagen, where overall

369

material toughness arises from high elasticity at low deformation forces and high stiffness at

370

high deformation forces.

371

Scaffold fabrication is a critical factor in whether or not the implant will facilitate

372

attachment and growth of new cells. To produce a PHBU scaffold with the necessary porous

373

architecture, two methods for pore production were utilized. It was anticipated that the smaller

374

pore size, produced by thermally induced phase separation, would enhance cellular attachment

375

and nutrient transport, while the larger pore size from salt leaching could promote

376

vascularization of the scaffold if necessary. While there is agreement that high porosity is

377

necessary for nutrient transport and cell attachment35, there are conflicting arguments on the

378

size of pore necessary for the most efficient cell adhesion to the scaffold. Certain studies have

379

shown success with scaffolds containing 100 µm are necessary for cell migration and transport,

383

and pore sizes >300 µm are preferred for the growth of capillaries and new bone.6 These

384

previous studies guided the approach taken in our study where PHBU scaffolds with two

385

different pore sizes were produced: small pore sizes were generated through thermally induced

386

phase separation of the polymer in frozen 1,4-dioxane, and large pore sizes were introduced

387

through a salt-leaching technique with sodium chloride. The benefit of the scaffold fabrication

388

method described here is the ability to alter pore sizes as necessary for the desired application

389

with minor procedural changes. Tissue-engineering scaffolds that require larger vasculature or

390

bone formation need only exchange the sodium chloride crystals for a larger porogen. Based on

391

the results shown here, scaffolds with small 30 µm pores promoted higher extracellular matrix

392

formation, which may be due to better cell adhesion. The reduced ECM formation observed on

393

PHBU scaffolds with only large pores (>300 µm) is likely due to surface topology, and not 15



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Page 16 of 36

394

nutrient transport due to the localization of the cells to the surface during the short incubation

395

period.

396

Cross-linking the PHBU scaffold resulted in a dramatic increase in tensile strength. With

397

tensile strengths as high as 26.2 MPa, this material may find use as cartilage or other

398

connective tissue substitute3,46. Interestingly, the cross-linked scaffolds were able to maintain

399

their geometry better than the non-cross-linked scaffolds after incubation for 48 h in hMSC

400

expansion media at 37°C. When physically transferring the cell-seeded scaffolds for

401

microscopy, non-cross-linked PHBU scaffolds were easily damaged despite gentle handling,

402

whereas the cross-linked PHBU scaffolds had higher durability and easily maintained their

403

shape even with rough handling. Cross-linking drastically improved the performance of the

404

scaffold in an aqueous environment, which is a notable drawback to other biopolymers used as

405

tissue-engineering materials47. This aqueous stability may become important for maintaining the

406

strength of the scaffold after extended exposure in vivo.

407

An important point of consideration is the effect of cross-linking on the rate of

408

biodegradation.

It

is

known

that

cross-linking

can

affect

the

rate

of

polymer

409

biodegradation,46,47,48,49 however these differences vary drastically depending on the type of

410

polymer and degree of cross-linking. In these studies, biodegradation took up to twice as long

411

(16 vs 8 days) in highly cross-linked samples. Further study into the biodegradation of cross-

412

linked PHAs would be necessary to determine whether their rates are suitable for tissue-

413

engineering scaffolds.

414

This study has shown that thiol-ene click chemistry can be used to efficiently cross-link

415

unsaturated PHA polymers in order to dramatically improve the strength of the materials. It was

416

found that the chemical cross-linking of the PHA scaffolds did not result in significant cytotoxicity

417

toward human mesenchymal stem cells. While it is encouraging that the new chemical cross-

418

links are not cytotoxic, further tests are necessary to conclude that the cross-linked material

16


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419

remains biocompatible.

In future in vivo studies, biocompatibility could be assessed by

420

implanting the cross-linked PHA into mice and monitoring inflammation.

421

Thiol-ene click chemistry is not limited to cross-linking reactions. This chemistry may

422

also be employed to incorporate new functional groups into the side chain of the polymer,

423

leading to control over other properties such as polarity, solubility, processing temperatures,

424

crystallinity, and biodegradability dependent on the modification. By controlling the repeating

425

unit composition of the PHA polymers produced by E. coli LSBJ, we can now approach a 2-

426

dimensional design framework where the polymer’s properties can be defined by the ratio of

427

SCL to MCL repeating units in addition to the new chemical moieties attached to the side chains

428

resulting in a matrix of tunable properties.

429

Conclusions

430

The three objectives of this work were to 1) produce an unsaturated PHA that could be

431

cross-linked by thiol-ene click chemistry, 2) determine the change in the physical properties of

432

PHA based on cross-link density, and 3) assess the viability of human cell cultures in the

433

presence of the chemically modified PHA. This work was successful in producing a SCL/MCL

434

PHA copolymer with tunable repeating unit composition for chemical modification. Using thiol-

435

ene click chemistry, cross-linked PHBU samples were produced with varying cross-link density

436

based on the concentration of the cross-linking reagent. The change in physical properties of

437

the highly cross-linked PHBU to 26.2 MPa tensile strength and 205.5-280.5 MPa Young’s

438

modulus will allow for new soft tissue applications unavailable to unmodified PHBU. When

439

cultured with hMSC, PHBU and cross-linked PHBU did not have a significant impact on cell

440

viability. For use in tissue engineering, it was critical that the chemical modification by thiol-ene

441

click chemistry not induce cytotoxicity, which would have rendered the scaffolds useless for

442

regenerative medicine applications. This study represents a critical first step in the use of

443

chemically modified PHA in tissue-engineering scaffolds. The lack of cytotoxicity toward hMSC

17



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Page 18 of 36

444

will allow for future experiments to assess the in vivo biocompatibility and biodegradation of

445

chemically cross-linked PHA. The three objectives of this work were accomplished with the

446

production of an unsaturated PHA copolymer with improved physical properties through

447

chemical cross-linking that did not adversely affect cellular viability for tissue engineering

448

applications.

449 450

Author Information

451

Corresponding Author

452

*E-mail: [email protected]; Telephone: 315-470-6854; Fax: 315-470-6856.

453

Author Contributions

454

The manuscript was written by Alex C. Levine and Christopher T. Nomura. Experiments were

455

designed by Alex C. Levine, Christopher T. Nomura, and Keiji Numata. Angelina Sparano and

456

Frederick F. Twigg assisted in PHA production. Alex C. Levine carried out all other experiments.

457

Acknowledgements

458

The authors acknowledge support from NSF CBET 1263905 awarded to C.T. Nomura and an

459

award to A.C. Levine from the NSF East Asia Pacific Summer Institutes (EAPSI) program in

460

collaboration with the Japanese Society for the Promotion of Science (JSPS). Special thanks

461

go to Keiji Numata, who was willing to host this research project in his lab at the RIKEN. We

462

also thank Jo-Ann Chuah and Jose Manuel Ageitos for their assistance with hMSC cultures and

463

microscopy. F.F.Twigg was supported by NSF REU 1156942 awarded to J.M. Hasenwinkel

464

and P.T. Mather of the Syracuse Biomaterials Institute.

465 466 467 468 469 470

18


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620 621 622

(51)

Goissis, G.; Junior, E. M.; Adriana, Rosemary Marcanto, C.; Carlos, R.; Lia, C.; Cancian, D. C. J.; Carvalho, W. M. De. Biocompatibility studies of anionic collagen membranes with different degree of glutaraldehyde cross-linking. Biomaterials 1999, 20, 27–34.

623 624 625 626 627 628 629 630 631 632 633 634 635 636 637 638 639 640 641 642 643 644 645 646 647 648 649 650 651 652 653 654 655 656 657 658 659 660 661 662 663 664 665 23



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

666 667 668 669 670

671 672 673 674 675 676 677 678 679 680 681 682 683 684 685 686 687 688 689 690 691 692 693 694 695 696 697 698 699 700 701 702 703 704 705 706 707 708 709 710 711 712 713

Table 1: Molecular weight and yield of PHBU samples a PHBU Mn Mw Mw/Mn 5:95 71.7 ± 4.0 208.4 ± 7.3 2.9 ± 0.1 10:90

63.7 ± 2.1

188.4 ± 2.8

20:80

75.8 ± 2.8

25:75

67.2 ± 3.4

50:50

93.2 ± 2.0

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Yield (% CDW) 43.4 ± 3.9

3.0 ± 0.1

43.8 ± 2.3

209.1 ± 3.3

2.8 ± 0.1

N.D.

201.3 ± 5.0

3.0 ± 0.1

N.D.

227.1 ± 2.0

2.4 ± 0.1

N.D.

a

Ratio of 3-hydroxyundecenoyl to 3-hydroxybutyrl repeating units in the polymer. Mn, apparent number average molecular weight as compared to polystyrene standard; Mw, apparent weight average molecular weight as compared to polystyrene standard; Mw/Mn, polydispersity; CDW, cell dry weight; N.D., No data

24


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1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

714 715 716 717 718


Table 2: Physical properties of cross-linked PHBU Modulus Stress Strain Sample (MPa) (MPa) (%) PHBU

719 720 721

367.5±43.8

8.5±0.5

88.8±19.2

PHBUX50%

a

258.3±67.9

11.2±0.2

120.3±16.5

PHBUX98%

b

205.5±58.7

26.2±1.1

95.2±8.3

a

b

50%gel fraction; 98% gel fraction; Modulus, Young’s modulus; stress, tensile strength; strain, elongation at break; MPa, megapascals

722 723 724 725 726 727 728 729 730 731 732 733 734 735 736 737 738 739 740 741 742 743 744 745 746 747 748 25



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

749 750

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Table 3: Comparative properties of natural and synthetic polymers Modulus Stress Strain Reference (MPa) (MPa) (%) a 367.5 8.5 88.8 This study PHBU PHBUX98% 205.5 26.2 95.2 This study Polylactic acid 1200 28 6 41 Polycaprolactone 400 16 80 41 Skin 0.1-0.2 7.6 60-80 3,39 Fibrocartilage 159.1 27.5 3 Ligaments 303.0 29.5 3 Low density polyethylene 200 10 620 25 Polytetrafluoroethylene 500 27.5 3 Silicone rubber 8 7.6 3 a Polymer generated from 10-undecenoic acid and sodium butyrate by E.coli LSBJ Sample

751 752 753 754 755 756 757 758 759 760 761 762 763 764 765 766 767 768 769 770 771 772 773 774 775 776 777 778 779 780 781 782 783 784 785 786 787 788 789 790 791

26


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1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

792 793 794 795 796 797 798 799 800 801 802 803 804 805 806 807 808 809 810 811 812 813 814 815 816 817 818 819 820


For Table of Contents Use Only:

Influence of cross-linking on the physical properties and cytotoxicity of polyhydroxyalkanoate (PHA) scaffolds for tissue engineering Alex C. Levine1, Angelina Sparano1, Frederick F. Twigg1, Keiji Numata2, Christopher T. Nomura1,3* 1

Department of Chemistry, State University of New York − College of Environmental Science and Forestry, 1 Forestry Drive, Syracuse, New York 13210, United States 2 Enzyme Research Team, RIKEN, 2-1 Hirosawa, Wako, Saitama 351-0198, Japan 3 Center for Applied Microbiology, State University of New York − College of Environmental Science and Forestry, 1 Forestry Drive, Syracuse, New York 13210, United States

Synopsis In this study, an engineered strain of Escherichia coli was used to produced an unsaturated polyhydroxyalkanoate polymer for use in tissue-engineering scaffolds. Produced from an unnatural fatty acid which allowed for chemical modification, the polymer was cross-linked using thiol-ene click chemistry and assessed for changes in physical properties and cytotoxicity. Results showed that cross-linking the polymer increased tensile strength by more than 3 fold, and that this chemical modification was not significantly toxic toward human mesenchymal stem cells. Scaffolds fabricated from this cross-linked biopolyester will be relevant for soft tissue replacement such as cartilage. This work shows that polymers produced from renewable sources can be used in tissue-engineering materials, allowing for their sustainable production from carbon sources such as fatty acids.

UV PETMP

Scaffold Fabrication

821

Tissue Engineering Scaffolds

27



A

Polymer Mole Ratio

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

0.3

B

0.25 0.2 0.15 y = 0.054ln(x) + 0.274 R² = 0.986

0.1

0.05 0

0

0.2

0.4

0.6

0.8

1

Substrate Mole Ratio Figure 1: Analysis of produced poly[(R)-3-hydroxybutyrate-co-(R)-3-hydroxy-10-undecenoate] (PHBU). A) Standard curve of PHBU. The 10-undecenoate content in the copolymer was monitored relative to the amount of 10-undecenoic acid present in the culture medium. The 10-undecenoate content in the polymer ranged from 5-30% with medium 10-undecenoate compositions ranging from 2-50%. The trend of unsaturated fatty acid uptake was non-linear, with less than incorporation after 12% PHU content. B) 1 HNMR of PHBU copolymer with 7% PHU content. Composition was determined using the ratio of the alkene proton signal at 5.8 ppm to the polymer backbone stereocenter proton at 5.3 ppm.


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1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Figure 2: Scanning electron micrographs of porous PHA scaffolds. Large pore sizes (>200 µm) are the result of salt leaching (A), whereas small pore sizes are fabricated using thermally induced phase separation (B). The combination of large and small pores should facilitate in-growth of cells and nutrient transport, respectively.



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

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PHBU PHBUX50% PHBUX98%

1660

1650

1640

1630

1620

Wavenumber (cm-1)

3100

2600

PHBU PHBUX50% PHBUX98%

2100

1600

1100

600

Wavenumber (cm -1) Figure 3: FT-IR of PHBU and cross-linked PHBU polymers. Evidence of successful cross-linking was -1 seen in the reduction of the signal for alkene functional group at 1641 cm . Extent of modification correlated to the concentration of cross-linking agent in the reaction.

A

B ACS Paragon Plus Environment

Page 31 of 36

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Figure 4: Fluorescence microscopy of human mesenchymal stem cells (hMSC) grown on cross-linked PHBU scaffolds for 48 hours. Live cells were stained with calcein AM, a membrane permeable dye that fluoresces green when exposed to intracellular esterases, and dead cells were simultaneously stained with ethidium homodimer-1, a membrane impermeable die that fluoresces red when bound to nucleic acids. A) Cells seeded on PHBU scaffold. B) Cells seeded on PHBUX98% scaffold. Cell viability was unaffected by cross-linking.



1.6

Absorbance 490 nm

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

1.4

1.2 1 0.8

0.6 0.4 0.2

0 PHBU

PHBUX50% + control

- control

Figure 5: MTS viability assay of hMSC grown on PHBUX50% scaffold. The differences between the positive control and PHBUX50% scaffolds, and between the PHBU and PHBUX50% scaffolds were statistically insignificant as determined by unpaired T-test values of p=0.82 and p=0.54, respectively.


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1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Figure 6: Scanning electron micrographs of human mesenchymal stem cells grown on and then fixed to cross-linked PHA scaffolds with large (>200 µm) pore size (A and B) small (30 µm) pore size (C and D). Greater amounts of extracellular matrix material was observed on regions of scaffold consisting mainly of small pore size.



30

A Stress (MPa)

25

20 15 10 5 0 0

25

50 75 Strain (%)

100

0

25

50 75 Strain (%)

100

30

B

25

Stress (MPa)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

20 15

10 5 0

Figure 7: Stress-strain curves of A) PHBU and B) PHBUX98%. An increase in tensile strength from 8.5 MPa in the non-cross-linked PHBU to 26.2 MPa in the cross-linked PHBU was observed. The shape of the stress strain curve became nearly J-shaped in the cross-linked polymer, where increasing strain on the sample resulted in continued increase in observed stress.


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1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


Scheme 1: Reaction scheme of the thiol-ene cross-linking of PHBU.



1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

169x70mm (96 x 96 DPI)


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Influence of Cross-Linking on the Physical Properties and Cytotoxicity

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