**1. Introduction**

Periodontitis is a chronic inflammatory oral disease caused by bacteria infection [1]. Typically, the infection destroys periodontal cells including gingival fibroblast, periodontal ligament fibroblasts, and alveolar bone, which are the supporting tissue and bone that hold the tooth. As the disease progresses, more oral tissues are damaged, causing deep pockets, which eventually lead to teeth loss if left untreated [2].

Tissue engineering has been employed to regenerate the lost periodontal tissues and restore both structure and function. In this regard, three dimensional (3D) porous scaffolds represent important components for tissue engineering as a supporting material for cell proliferation or differentiation before being applied to repair the damaged area [3,4]. Scaffolds provide attachment sites and structural guidance for cells that enable them to synthesize appropriate extracellular matrix (ECM) proteins and ultimately proliferate into functional tissues [5]. In addition, the choice of scaffold can be critical as its chemical and

**Citation:** Phuegyod, S.; Pramual, S.; Wattanavichean, N.; Assawajaruwan, S.; Amornsakchai, T.; Sukho, P.; Svasti, J.; Surarit, R.; Niamsiri, N. Microbial Poly(hydroxybutyrate-cohydroxyvalerate) Scaffold for Periodontal Tissue Engineering. *Polymers* **2023**, *15*, 855. https:// doi.org/10.3390/polym15040855

Academic Editor: Shashi Kant Bhatia

Received: 4 January 2023 Revised: 31 January 2023 Accepted: 2 February 2023 Published: 9 February 2023

**Copyright:** © 2023 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

physical properties provide guidance cues for the cells to behave appropriately. Scaffold biomaterials for successful tooth regeneration applications should have some requirements such as being biocompatible, biodegradable, and possess mechanical properties that are consistent with the implanted area as well as being used in the appropriate amount and with an accessible volume of porosities for the diffusion of oxygen, cells, and nutrients [6,7]. To date, many polymeric materials have been reported to create biodegradable scaffolds for dental tissue engineering including poly(lactide) (PLA) [8], poly(lactide-co-glycolide) (PLGA) [9,10], and polycaprolactone (PCL) [11–13].

Polyhydroxyalkanoates (PHAs) are aliphatic polyesters synthesized by microorganisms to store excess carbon and energy. Poly(hydroxybutyrate-co-hydroxyvalerate) or P(HB-HV) copolymers are a member of the PHA family [14]. P(HB-HV) has shown great potential for tissue engineering with attractive characteristics of natural origin, biocompatibility, and biodegradability. The properties are adjustable by changing the content of the HV unit. P(HB-HV) are less crystalline, less stiff, and more flexible than the PHB homopolymer due to the incorporation of the HV monomer in the polymer chain [15]. Recently, studies have revealed that different types of scaffolds made with P(HB-HV) demonstrate desirable advantages for tissue engineering. The application of macroporous P(HB-8HV) matrices in the repair of full-thickness cartilage defects in rabbits in vivo was reported by Kose et al. At 8 and 20 weeks after seeding, in vivo results with chondrocyte seeded P(HB-8HV) matrices presented early cartilage formation resembling normal articular cartilage and revealed minimal foreign body reaction. This study also showed that P(HB-8HV) matrices maintained their integrity for 21 days and permitted appropriate gradual degradation and allowed for tissue remodeling to take place [16]. Abazari et al. demonstrated the increased survival rate and insulin-producing cell (IPC) differentiation potential of induced pluripotent stem cells (iPSCs) cultured on a nanofibrous 3D P(HB-5HV) scaffold in comparison with the 2D substrate. iPSCs-P(HB-5HV), as a promising cell-copolymer construct, could potentially be applied in pancreatic tissue engineering applications to diabetic patient treatment [17]. The P(HB-3HV) scaffold was tested for degradation in simulated body fluid (SBF), pH 7.4. After 8-week periods, the P(HB-3HV) scaffolds revealed about 51% weight loss along time due to the high porous structure when compared with the dense and compact films, which showed about a 9% weight loss. Culturing of MC3T3-E1 pre-osteoblast cells on the P(HB-3HV) scaffold samples obtained after 6 weeks of degradation did not lead to the formation of cytotoxic components [18]. In spite of extensive research on P(HB-HV) and their blends as scaffolds for tissue engineering, the HV molar contents of the available commercial P(HB-HV) published are 12 mol% or lower. P(HB-HV) films consisting of various HV content (5–80%) produced by *Paracoccus denitrificans* have been reported to be biocompatible with connective tissue, bone, and dermal fibroblast cells [19]. *Haloferax mediterranei* ES1 produced P(HB-HV) nanofibrous meshes were also shown to be excellent in vitro and to show in vivo biocompatibility with skin tissues [20]. As reported earlier, the more flexible P(HB-HV) with HV contents of 50 mol% can be successfully biosynthesized from bacteria *Cupriavidus necator* H16. This material has already been employed as a drug delivery platform [21–23]. Until now, there have been no studies available in the literature concerning the application of P(HB-50HV) produced by *C. necator* H16 as a scaffold to support cell growth and promote tissue regeneration.

In this study, the 3D porous scaffolds were fabricated from bacterial derived P(HB-50HV) via a particulate leaching method using salt particles as a strategy for the regeneration of periodontal cells. Comparisons were made with scaffold prepared from poly(hydroxybutyrate) (PHB) and poly(hydroxybutyrate-co-12%hydroxyvalerate) (P(HB-12HV)), and the well-established synthetic polycaprolactone (PCL). The scaffolds were characterized with respect to the morphology of the surface and cross section, porosity, mechanical strength, and protein absorption. Subsequently, biological performance of the scaffolds in terms of biocompatibility and cell proliferation was assessed. In this regard, human gingival fibroblasts (HGFs) [10] and periodontal ligament stem cells (PDLSCs) [24]

were used since they have been widely studied for the initial evaluation of biomaterials for periodontal tissue engineering applications.

#### **2. Materials and Methods**

#### *2.1. Materials*

Poly(hydroxybutyrate) (PHB, Mw 3.5 × 105 g/mol), poly(hydroxybutyrate-cohydroxyvalerate) containing HV content 12 mol% (P(HB-12HV), Mw 2.5 × 105 g/mol), polycaprolactone (PCL, Mw 6.5 × 104 g/mol), and 3-(4,5-Dimethyl-2-thiazolyl)-2,5-diphenyl-2H-tetrazolium bromide (MTT) were obtained from Sigma-Aldrich (St. Louis, MO, USA). Chloroform, methanol, and ethanol were purchased from RCI Labscan (Bangkok, Thailand). Sodium chloride (NaCl) (Ajax Chemicals Ltd., Sydney, Australia) with the particle size range of 425–500 μm was obtained by sieving through an analytical sieve shaker Octagon digital (Endecotts Ltd., London, UK) using two certified sieve sizes with 425 and 500 μm. Dulbecco's modified Eagle medium (DMEM), fetal bovine serum (FBS), and trypsin-EDTA were obtained from Invitrogen (Carlsbad, CA, USA). All chemicals and solutions were used as supplied without further purification. Poly(hydroxybutyrate-co-hydroxyvalerate) with 50% HV content (P(HB-50HV) Mw 1.69 × <sup>10</sup><sup>6</sup> g/mol) was biosynthesized by in house bacterial cultivation according to a previously described protocol [21].

#### *2.2. Characterization of Polymer Films*

Thin polymeric films of PHB, P(HB-12HV), P(HB-50HV), and PCL were prepared by a casting method using 10 mL of 2% (*w*/*v*) polymer stock solution in chloroform on a clean Petri dish. Chloroform was evaporated in a fume hood at room temperature for 24 h. The final thickness of film ranged from 0.05 to 0.10 mm.

The Fourier transform infrared (FTIR) spectra of all PHA and PCL thin films were obtained with a Perkin-Elmer FTIR ATR-FTIR spectrometer (Perkin-Elmer, Spectrum GX FTIR; Shelton, CT, USA). The sample spectra were recorded over 20 scans between 400 and 4000 cm−<sup>1</sup> wavenumbers at a resolution of 4 cm<sup>−</sup>1.

The hydrophilicity of the polymeric surfaces was examined by an optical bench-type contact angle goniometry DM-CE1 (Kyowa Interface Science, Niiza, Japan) using a sessile drop method at room temperature.

#### *2.3. Fabrication and Characterization of Scaffolds*

Salt-leached scaffolds of PHB, P(HB-12HV), P(HB-50HV), and PCL were fabricated following the established procedure [25]. In brief, the polymer was dissolved in chloroform to prepare a 5% (*w*/*v*) stock solution. The 1 mL polymer solution was then poured on a bed of sieved NaCl particles (with size range of 425–500 μm) in a clean glass vial. The weight ratio of porogen (NaCl) to polymer was set at 9:1. The scaffolds were placed in a fume hood at room temperature for the slow evaporation of chloroform over 2 days followed by repeated rinsing with distilled water to remove any residual salt and air-dried. All scaffolds were prepared as a cylindrical shape with 10 mm diameter and 3 mm height.

The fabricated scaffolds were mounted onto an aluminum stub, gold-coated, and then observed by scanning electron microscopy (SEM, JSM-6360; JEOL Techniques, Tokyo, Japan) with an accelerating voltage of 20 kV for the surface topography and cross section images.

The porosity or void volume fraction Vf (%) of the scaffold was calculated using the following equation:

$$\mathbf{V}\_{\mathbf{f}} = (1 - (\rho\_{\mathbf{s}}/\rho\_{\mathbf{m}})) \times 100$$

where ρ<sup>s</sup> is the apparent density of the porous scaffold and ρ<sup>m</sup> is the density of the polymer material [26].

#### *2.4. Compressive Mechanical Testing of Scaffolds*

The scaffolds were subjected to mechanical measurements under compressive mode in order to determine the compressive stress and compressive modulus (E). The tests were performed at room temperature using a Texture analyzer (TA-XT2i, Stable Micro Systems, Ltd., Godalming, UK) with a 50 kN load cell at a crosshead speed of 0.1 mm/s [27]. Cylindrical specimens were tested under both dry and wet conditions. The load deformation curves of the samples obtained were converted into stress–strain curves. The compressive stress (MPa) was used to calculate the secant modulus according to the following equation:
