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. 2021 Oct 25;6(43):28788–28803. doi: 10.1021/acsomega.1c03569

Development of an Antimicrobial-Coated Absorbable Monofilament Suture from a Medical-Grade Poly(l-lactide-co-ε-caprolactone) Copolymer

Montira Sriyai †,§, Jagkrit Tasati ‡, Robert Molloy §, Puttinan Meepowpan ‡,§, Runglawan Somsunan ‡,§, Patnarin Worajittiphon ‡,§, Donraporn Daranarong §,∥, Jomkwan Meerak §,⊥, Winita Punyodom ‡,§,*
PMCID: PMC8567407  PMID: 34746572

Abstract

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In this study, a medical-grade poly(l-lactide-co-ε-caprolactone) (PLC) copolymer with a monomer ratio of l-lactide (L) to ε-caprolactone (C) of 70:30 mol % for use as an absorbable surgical suture was synthesized via ring-opening polymerization (ROP) using a novel soluble liquid tin(II) n-butoxide (Sn(OnC4H9)2) as an initiator. In fiber fabrication, the process included copolymer melt extrusion with a minimal draw followed by sequential controlled hot-drawing and fixed-annealing steps to obtain oriented semicrystalline fibers with improved mechanical strength. For healing enhancement, the fiber was dip-coated with “levofloxacin” by adding the drug into a solution mixture of acetone, poly(ε-caprolactone) (PCL), and calcium stearate (CaSt) in the ratio of acetone/PCL/CaSt = 100:1% w/v:0.1% w/v. The tensile strength of the coated fiber was found to be increased to ∼400 MPa, which is comparable with that of commercial polydioxanone (PDS II) of a similar size. Finally, the efficiency of the drug-coated fiber regarding its controlled drug release and antimicrobial activity was investigated, and the results showed that the coated fiber was able to release the drug continuously for as long as 30 days. For fiber antimicrobial activity, it was found that a concentration of 1 mg/mL was sufficient to inhibit the growth of Staphylococcus aureus (MRSA), Escherichia coli O157:H7, and Pseudomonas aeruginosa, giving a clear inhibition zone range of 20–24 mm for 90 days. Cytotoxicity testing of the drug-coated fibers showed a %viability of more than 70%, indicating that they were nontoxic.

1. Introduction

In the last five decades, there has been a markedly increased interest in synthetic bioadsorbable polymers for medical devices such as surgical sutures, vascular scaffolds, nerve guides, plates and screws, drug-eluting stents, etc.1−6 Aliphatic polyesters such as poly(l-lactide) (PLL), poly(d-lactide) (PDL), poly(d,l-lactide) (PDLA), poly(rac-lactide), polyglycolide (PGA), poly(ε-caprolactone) (PCL), poly(δ-valerolactone) (PVL), and their other high-molecular-weight co- or terpolymers such as poly(l-lactide-co-ε-caprolactone) (PLC) and poly(l-lactide-co-ε-caprolactone-co-glycolide) (PLCG) were found to have a multitude of uses in medical applications.7−19 Typically, biodegradable polyesters can be obtained via ring-opening polymerization (ROP) in the bulk.20 For the ROP of lactides and lactones, tin(II) 2-ethylhexanoate or stannous octoate (Sn(Oct)2) has been approved for use as a food additive by the US FDA and is the most common initiator used due to its effectiveness and versatility. It is also easy to handle and soluble in most common organic solvents and monomers.21−23 However, to improve its effectiveness, Sn(Oct)2 is usually reacted with an alcohol coinitiator (ROH) via a coordination-insertion mechanism, resulting in the formation of tin(II) monoalkoxide ((Oct)Sn(OR)) and tin(II) dialkoxide (Sn(OR)2), which then become the true initiators.24−27 Therefore, in this work, a novel liquid tin(II) n-butoxide (Sn(OnC4H9)2) was synthesized for use as an initiator in the ROP synthesis of medical-grade poly(l-lactide-co-ε-caprolactone).28−31

An absorbable surgical suture acts as a temporary scaffold while the process of wound healing is taking place following surgery. The suture is used to hold tissues together until sufficient collagen synthesis has taken place to hold the wound together unassisted.32,33 An absorbable suture can be fabricated as multifilaments or monofilaments. Whereas multifilament sutures are made up of tiny fibers braided or twisted together, monofilament sutures are single-stranded fibers that have no braiding or twisting. Both have their advantages and disadvantages. Multifilament sutures are soft, flexible, and easy to handle, but there are interstices formed by the relatively loose braid of the fibers that permit serum and blood to penetrate the suture and form a perfect refuge for bacteria. Therefore, multifilament sutures should not be used in a wound that is likely to become infected. Moreover, their outer surface is rough, which can make it difficult to pull the suture through delicate tissue without tearing the tissue in the process. Monofilament sutures have less affinity for bacteria and a smoother surface, which are major advantages in exhibiting less tissue drag. On the other hand, they are quite stiff and difficult to handle and tie, which are their disadvantages.34−36

For absorbable surgical suture production, the range of polymer options is mainly restricted to aliphatic polyesters and co- or terpolyesters. This is because these polymers have been shown to possess the best balance of properties required in suture applications such as good handling properties, high knot security, and minimal or no inflammatory reactions.37−40 In general, an absorbable suture loses most of its strength and degrades within 60 days by hydrolysis, giving carbon dioxide and water byproducts, which can be excreted from the human body.41,42

Nowadays, absorbable sutures in the market are Dexon (polyglycolic acid), Vicryl (polyglactin 910), Monocryl (polyglecaprone 25), Maxon (polyglyconate), PDS II (PDO, polydioxanone), Surgisorb M (PLC, 75/25 poly(l-lactide-co-ε-caprolactone)), and MonoMax (PHB, poly-4-hydroxybutyrate), which possess 40–60, 60–90, 90–120, 180, 200, and 400 days of complete absorption, respectively.43−47 Among these sutures, PDO, PLC, and PHB are classified as long-term absorbable monofilament sutures. Although PLC copolymers (Figure 1) have been reported in the literature for use in a wide range of biomedical applications such as drug delivery systems, barrier membranes, absorbable nerve guides, bioadhesives, and scaffolds for tissue engineering, relatively little attention has been paid to absorbable sutures, possibly because of their slow rate of absorption (typically > 12 months) in the human body.48−58 However, for use with the specific tissues or organs such as tendon and ligament repairs or abdominal wall closure, where slow absorption is required, PLC copolymers have been shown to be effective and versatile materials that can be tailored through their composition to meet the specific property requirements of a given application.59−63

Figure 1.

Figure 1

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Synthesis of a poly(l-lactide-co-ε-caprolactone) (PLC) random copolymer.

Typically, the most common adverse effect when using surgical sutures is the microbial colonization of suture material in operative wounds (surgical site infection, SSI).64,65 Therefore, to inhibit microbial colonization of suture material in operative wounds, Triclosan (TC)-coated suture materials with antibacterial activity have been developed and widely used.66 TC is a broad-spectrum phenol family antiseptic, used for a long time as a safe and effective antimicrobial agent, against the most common bacterial pathogens that cause SSI such as Staphylococcus aureus and Escherichia coli. The antimicrobial efficacy of this material in reducing both bacterial adherence to the suture and microbial viability has been both proven in vitro and in vivo.67−70 However, only recently, has it been reported that TC commonly used in commercially available sutures is suspected carcinogenic.71 Consequently, levofloxacin (LVFX, C18H20FN3O4, (S)-9-fluoro-2,3-dihydro-3-methyl-10-(4-methylpiperazin-1-yl)-7-oxo-7H-pyrido[1,2,3-de]-1,4-benzoxazine-6-carboxylic acid) (Figure 2) is another choice for use as a replacement for the TC antimicrobial drug to enhance the SSI inhibition. Nonetheless, presently, only a few investigations on the efficiency of LVFX-coated PLC sutures for reducing SSI have been reported.72−75 Therefore, this paper now describes the synthesis, characterization, fiber fabrication and mechanical strength improvement, and LVFX coating of a synthesized medical-grade PLC copolymer for potential use as a long-term absorbable monofilament surgical suture.

Figure 2.

Figure 2

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Chemical structure of antibiotic drug levofloxacin (LVFX).

2. Results and Discussion

2.1. Synthesis of the Liquid Sn(OnC4H9)2 Initiator

The synthesized liquid tin(II) n-butoxide, Sn(OnC4H9)2, was obtained as a viscous, dark yellow liquid of approximately 70–75% yield. Its physical appearance, solubility test results, and structural confirmations by various methods were described elsewhere.76

2.2. Synthesis of the Medical-Grade PLC Copolymer

In this work, medical-grade poly(l-lactide-co-ε-caprolactone) of LL/CL composition = 70:30 mol % for use as an absorbable monofilament suture was synthesized via bulk ring-opening polymerization (ROP) using liquid Sn(OnC4H9)2 as an initiator at 120 °C for 96 h. After being cut, ground, and purified, the PLC copolymer was obtained as white solid granules with >80% yield.

For the structural characterization of the synthesized PLC copolymer, the typical 1H NMR spectrum of PLC in deuterated chloroform at room temperature including the peak assignments for the various protons of the polymer structure is shown in Figure 3 and the chemical shifts (δ) are given in Table 1.

Figure 3.

Figure 3

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400 MHz 1H NMR spectrum of a medical-grade PLC synthesized using 0.01 mol % liquid Sn(OnC4H9)2 as an initiator at 120 °C for 96 h.

Table 1. Proton Assignments and Chemical Shifts of the 1H NMR Spectrum of PLC Synthesized Using 0.01 mol % Liquid Sn(OnC4H9)2 as an Initiator at 120 °C for 96 h.

proton assignments chemical shift range (δ, ppm)
b 5.11–5.17
c 4.04–4.17
g 2.29–2.32
a, d, f 1.51–1.59
e 1.38–1.40
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Table 2 shows the average molecular weights (M̅n and M̅w) and polydispersity index (Đ = M̅w/M̅n) of the PLC copolymer determined by gel permeation chromatography (GPC) using tetrahydrofuran (THF) as a solvent at 40 °C (flow rate = 1 mL/min) and calibrated with polystyrene (PS) standards. From the GPC results (see the Supporting Information), the obtained copolymer showed considerably high molecular weights with narrow polydispersity. Generally, melt extrusion of aliphatic polyesters requires a value of M̅n ≥ 3.0 × 104 Da, and so, the value in Table 2 is considered to be high enough for fiber fabrication.

Table 2. Molecular Weights and Dispersity of the Synthesized PLC Copolymer by GPC.

  average molecular weighta
  
copolymer/initiator number-average molecular weight, M̅n (Da) weight-average molecular weight, M̅n (Da) dispersity, Đb
PLC/0.01 mol % Sn(OnC4H9)2 7.03 × 104 1.23 × 105 1.76
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a

Average molecular weights as determined by GPC using THF as a solvent at 40 °C, flow rate = 1 mL/min, and polystyrene (PS) standards.

b

Đ = dispersity = M̅w/M̅n.

According to the ASTM F1925-17 requirements, characterization and analysis of the copolymer product by various techniques were obtained, as summarized in Table 3. From the results obtained, the synthesized PLC obtained was found to meet all of the requirements according to the standard specifications; therefore, it can be further used for fabrication into a monofilament suture. Moreover, it can be seen from Table 3 that the properties obtained from the synthesized PLC were close to those of the commercial Evonik’s Resomer LC 703 S of similar medical grade. However, it was found that the synthesized PLC contained more residual catalyst (Sn) than the commercial PLC and showed slightly lower inherent viscosity.

Table 3. Requirements for the Medical-Grade Poly(l-lactide-co-ε-caprolactone) (PLC) Copolymer in Accordance with ASTM F1925-17.

specification ASTM F1925-17a synthesized 70:30 PLCb commercial resomer LC 703 Sc
physical appearance (optional) white granule/flake white granule white granule
residual monomer LL + CL (%) <2 (total) 1.15 ± 0.01d (0.15% LL) (0.00% CL) 1.27 ± 0.02d (0.17% LL) (0.10% CL)
residual solvent (ppm) <1000 <1d <1d
residual catalyst as Sn (ppm) (optional) ≤150 60e 28e
residual heavy metal as Pb (ppm) ≤10 <1e < 1e
moisture content (%) (optional) ≤0.5 0.07f 0.21f
LL/CL composition variation from target ratio (% molar ratio) (optional) ±3 1g (69:31) 1g (69:31)
specific optical rotation, [α]D25 (deg) (optional)   118.41 ± 0.13h 115.13 ± 0.56h
inherent viscosity (dL/g) (optional) 1.3–1.8i 1.58i 1.61i
glass transition temperature, Tg (°C) (optional)   28.40j 32.60j
melting temperature, Tm (°C) (optional)   160.4j 162.3j
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a

ASTM F1925-17: standard specification for semi-crystalline poly(lactide) polymer and copolymer resins for surgical implants.

b

70:30 PLC synthesized via ROP in the bulk using 0. 01 mol % liquid Sn(OnC4H9)2 as an initiator at 120 °C for 96 h.

c

Commercial Evonik’s Resomer LC 703 S (medical-grade 70/30) poly(l-lactide-co-ε-caprolactone).

d

Residual monomers LL and CL (%) and solvents (ppm) as determined by gas chromatography with a flame ionization detector (GC-FID) using methylene chloride as a solvent and 2,6-dimethyl-α-pyrone as an internal standard.

e

Residual catalyst (Sn) and heavy metal (Pb) as detected using inductively coupled plasma-optical emission spectroscopy (ICP-OES).

f

Moisture content as determined by the Karl Fischer titration technique at 150 °C.

g

Chemical composition of copolymer PLC as calculated from the 1H NMR spectra by taking the ratio of the peak areas corresponding to the LL methine protons at δ = 5.0–5.3 ppm and the CL ε-methylene protons at δ = 3.9–4.2 ppm using a 400 MHz Bruker Avance II NMR spectrometer with CDCl3 as a solvent.

h

Specific rotation, [α]D25, of the commercial Evonik’s Resomer LC 703 S (medical-grade 70/30 poly(l-lactide-co-ε-caprolactone)) and the synthesized PLC measured by an ATAGO AP-300 Automatic Polarimeter at λ = 589 nm.

i

Inherent viscosity, ηinh, as measured from a 0.5% w/v solution in CHCl3 at 30 ± 0.1°C.

j

Tg and Tm of the commercial Evonik’s Resomer LC 703 S (medical-grade 70/30 poly(l-lactide-co-ε-caprolactone)) and the synthesized PLC measured by differential scanning calorimetry (DSC) (PerkinElmer DSC7 with Pyris 1 Software) under a flowing N2 atmosphere (20 mL/min) at a heating rate of 10 °C/min over the temperature range of 0–200°C.

2.3. PLC Fiber Fabrication and Mechanical Strength Improvement

The PLC copolymers were melt-extruded into monofilament fibers using a single-screw melt extruder with stretching units and associated equipment. Melt extrusion for fiber-forming, hot-drawing, and fixed-annealing for enhancing the fiber’s properties were optimized to find the best possible set of fiber processing conditions. The as-spun fibers obtained were of uniform diameter in the range of 0.9–1.2 mm and were clear and colorless with a smooth surface. On their emergence from the die, the as-spun fibers were cooled in a water bath with only minimal draw so that they would be largely amorphous and unoriented. The rather poor mechanical properties of the largely amorphous as-spun fiber were then improved by hot-drawing and fixed-annealing steps under well-controlled conditions. First, the as-spun fiber was hot-drawn at 60 °C at a high draw ratio (1:5) followed by fixed-annealing at 80 °C for 5 h. These hot-drawing and fixed-annealing steps were repeated once more using a 1:2 draw ratio at 90 °C (2nd hot-drawing) and 80 °C for 24 h (2nd fixed-annealing) to develop further the fiber’s oriented semicrystalline morphology and thereby increase its tensile strength. At last, the free-annealing process at 60 °C was carried out to stabilize the fiber morphology as a result of molecular relaxation, leading to fiber shrinkage.

2.4. Tensile Testing

The tensile property testing results and stress–strain curves of the PLC monofilament fibers at various stages of their processing are shown in Table 4 and Figure 4A. It can be seen that the as-spun fiber was very weak (stress at break ∼10 MPa) due to its largely amorphous nature with low levels of % crystallinity and molecular orientation. However, after the as-spun fiber was first hot-drawn, its strength increased by 850% following which it became more flexible after fixed-annealing at 80 °C for 5 h. A further second hot-drawing step induced more crystallinity and orientation and resulted in the highest strength, as shown in Figure 4A. Finally, to complete the processing operation, thermal treatment by a combination of fixed- and free-annealing was necessary to stabilize the fiber morphology. The fixed-length annealing was performed at 80 °C for 24 h to allow molecular relaxation to occur. This was then followed by free-annealing at 60 °C for a further 72 h to complete the stabilization process. It was found that stabilization could reduce fiber shrinkage during sterilization and storage to not more than 12%. In addition, Figure 4B shows the scanning electron microscopy (SEM) micrographs of the PLC monofilament fiber at the various stages of its processing.

Table 4. Diameter and Mechanical Properties of the PLC Monofilament Fiber at the Various Stages of Processing.

fiber at the various stages of processing diameter, avg ± SD (mm) (spec. USP 3–0 = 0.255–0.339) draw ratio (λ) stress at break, avg ± SD (MPa) strain at break, avg ± SD (%) initial modulus, avg ± SD (MPa)
as-spun 1.156 ± 0.01   17.60 ± 0.11 >1500 231.9 ± 25.2
1st hot-drawn 0.477 ± 0.01 5 149.2 ± 9.45 218.6 ± 0.46 892.0 ± 50.7
2nd hot-drawn 0.301 ± 0.01 2 434.5 ± 13.6 65.23 ± 13.6 2235 ± 32.1
fix-annealed 0.323 ± 0.01   375.7 ± 8.49 71.30 ± 19.4 1790 ± 59.9
free-annealed (final) 0.335 ± 0.01   330.3 ± 12.3 126.6 ± 9.30 1236 ± 32.9
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Figure 4.

Figure 4

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(A) Stress–strain curves and (B) SEM micrographs (magnification ×100) of the PLC monofilament fiber at the various stages of processing: (purple line) as-spun, (red line) 1st hot-drawn, (blue line) 2nd hot-drawn, (yellow line) fix-annealed, and (green line) free-annealed (final).

From the results obtained, the effects of the successive hot-drawing and fixed-annealing steps on the stress–strain curve of the PLC fiber are shown in Figure 8. It was found that the 1st hot-drawing at 50 °C at fast drawing rates, as employed here, forces the molecules to align along the fiber axis against the constraints of molecular entanglements and with minimal time for molecular relaxation. This inevitably creates strain within the matrix. Fixed-length annealing at an appropriate temperature above Tg (80 °C) then allows the molecules to rearrange themselves into less-strained conformations but without accompanying fiber shrinkage. This molecular relaxation has the combined effects of lowering both the tensile strength and initial modulus (i.e., stiffness). In the later stage, the 2nd hot-drawing step stretches the relaxed fiber, further orienting the molecules in both the amorphous and crystalline phases, thereby increasing the tensile strength and modulus. Finally, the free-annealing step at 60 °C for further 72 h was performed to complete the stabilization process. It was found that stabilization could reduce fiber shrinkage during sterilization and storage to not more than 12%.

Figure 8.

Figure 8

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Comparison of the weight loss profiles of (blue bar) uncoated; (yellow bar) 0.5, (green bar) 1, (purple bar) 5, and (grey bar) 10 mg/mL LVFX-coated; and (orange bar) PDS II monofilament fibers immersed in PBS, pH 7.40, at 50 °C.

2.5. Antimicrobial Drug-Coated PLC Monofilament Fiber

Figure 5 shows the stress–strain curves of the uncoated and LVFX-coated fibers with various concentrations and mechanical property values shown in Table 5. From the results obtained, there was a significant increase in the tensile strength of the coated sutures and mechanical properties increase with increasing the drug content. Therefore, the highest straight-pull tensile strength (∼400 MPa) was found in the PLC suture coated with 10 mg/mL LVFX, while the lowest value (336 MPa) was obtained from the 0.5 mg/mL LVFX-coated fiber. This strength increase might be due to the PCL in the coating formulation acting as an external lubricant by lowering the surface friction of the coated sutures around the bollard grips during testing. Accordingly, as the tensile strength and initial modulus of each fiber increased, a decrease in the fiber flexibility was observed.

Figure 5.

Figure 5

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Stress–strain curves of (blue line) uncoated and LVFX-coated PLC fibers with various concentrations of (yellow line) 0.5, (green line) 1, (purple line) 5, and (orange line) 10 mg/mL.

Table 5. Tensile Properties of the Uncoated and LVFX-Coated Fibers with Various Concentrations.

fiber diameter, avg ± SD (mm) stress at break, avg ± SD (MPa) strain at break, avg ± SD (%) initial modulus, avg ± SD (MPa)
uncoated 0.335 ± 0.01 330.3 ± 12.3 127.9 ± 9.30 1236 ± 32.9
0.5 mg/mL LVFX-coated 0.336 ± 0.01 336.0 ± 14.2 69.66 ± 4.87 1203 ± 23.6
1 mg/mL LVFX-coated 0.336 ± 0.01 341.2 ± 12.0 70.48 ± 5.43 1506 ± 65.4
5 mg/mL LVFX-coated 0.337 ± 0.00 355.6 ± 5.12 72.96 ± 5.69 1609 ± 52.0
10 mg/mL LVFX-coated 0.338 ± 0.00 398.7 ± 3.26 81.43 ± 5.62 1919 ± 31.7
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2.6. Confirmation of LVFX-Coated onto PLC Fibers

From the results obtained, SEM images and energy-dispersive X-ray fluorescence (EDXRF) elemental analysis data confirm the presence of carbon (C), oxygen (O), and fluorine (F), as shown in Table 6. Therefore, LVFX coating onto the fiber by the dip coating technique was successful. Moreover, Fourier transform infrared (FTIR) spectroscopy spectra of the PLC and LVFX-coated PLC fibers, as shown in Figure 6A,B, were obtained using a Nicolet iS 5 FTIR spectrometer (Thermo Fisher Scientific Inc.). Attenuated total reflectance (ATR) mode was used to record spectra with 32 scans over the range 400–4000 cm–1 at a resolution of 2 cm–1, giving the vibrational assignment data shown in Table 7.

Table 6. SEM Images and EDXRF Elemental Analysis Data of the LVFX-Coated Sutures with Various Concentrations.

2.6.

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Figure 6.

Figure 6

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FTIR spectra in the region between 4000 and 400 cm–1 of (A) PLC and (B) 10 mg/mL LVFX-coated PLC fibers.

Table 7. Vibrational Assignments of PLC and LVFX-Coated PLC Fibers from ATR-FTIR Spectroscopy.

PLC
 LVFX-coated PLC
vibrational assignments wavenumber (cm–1) vibrational assignments wavenumber (cm–1)
O–H stretching (chain end)   O–H stretching  
C–H stretching in CH, CH2, CH3 2944 C–H stretching (aromatic) 2950
C=O stretching 1754 C=O stretching 1723
C–H bending in CH, CH2, CH3 1455, 1384 C–N stretching (amine) 1294
C–O stretching (acyl oxygen) 1204, 1182 C–F stretching 1089
C–O stretching (alkyl oxygen) 1089    
CH2 bending 755    
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From Figure 6A,B and Table 7, though there is some overlap between functional groups of the uncoated and those of the LVFX-coated PLC, the C–N and C–F stretching bands at 1294 and 1089 cm–1 (Figure 6B) can be used to distinguish structural groups of the PLC associated with LVFX from the uncoated PLC.

2.7. In Vitro Drug Release Behavior

For in vitro drug release behavior study, the uncoated and LVFX-coated fibers with various drug concentrations of 0.5, 1, 5, and 10 mg/mL was carried out in phosphate-buffered saline (PBS) at 37 ± 1.0 °C for a period of 0, 1, 2, 4, 6, 8, 12, and 24 h and 0, 7, 14, 21, 28, and 30 days. The drug release profiles as a function of time intervals 0–24 and 0–720 h are shown in Figure 7A,B. It can be seen that the drug release profiles show a directly proportional relationship with an increased amount of the drug. From Figure 7B, the initial “burst” release was observed during the first 24 h for all concentrations of LVFX-coated fibers. After that, the release rate of all fiber samples remained rather stable.

Figure 7.

Figure 7

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In vitro drug release profiles of LVFX-coated PLC fibers with various concentrations in PBS pH 7.40 at 37 °C for (A) 0–24 and (B) 0–720 h.

2.8. Accelerated In Vitro Hydrolytic Degradation

2.8.1. Weight Loss Profile

The in vitro hydrolytic biodegradation rates at various immersion times of the uncoated, LVFX-coated with different concentrations, and the commercial PDS II fibers were monitored by measuring the weight loss of the samples, and the corresponding weight loss profiles are plotted in Figure 8. From the results obtained, it can be seen that there are no significant differences in the weight loss profile of the uncoated and fibers coated with LVFX. Moreover, the amount of LVFX used has no substantial effect on the weight loss profile. However, compared to that of PDS II, it was found that the weight loss of the uncoated and coated PLC fibers was relatively slow with approximately 10% weight remaining after 30 days of accelerated biodegradation. This is probably due to the hydrophobic ε-caprolactone component (30 mol %) making hydrolysis in the PBS medium more difficult, thereby resulting in a lower absorption rate.

2.8.2. Mechanical Strength

Figures 9 and 10 show the comparison of the obtained % breaking strength retention (BSR) and stress at break (MPa) values of the uncoated, LVFX-coated, and commercial PDS II at the time of in vitro degradation from day 1 to day 30. Among the uncoated and coated PLC samples, from day 0 to day 7, the fastest strength decrease rate was found in the 10 mg/mL LVFX-coated fibers. It was also found that on day 3, all of the PLC fibers lost their strengths by 50% and were completely absorbed within 28 days. However, compared with PDS II after immersion in PBS for 0–35 days, PDS II showed a much faster rate of strength reduction at day 7 and completely lost its total strength within 14 days.

Figure 9.

Figure 9

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Comparison of % BSR vs time profiles for (blue bar) uncoated; (yellow bar) 0.5, (green bar) 1, (purple bar) 5, and (grey bar) 10 mg/mL LVFX-coated PLC; and (orange bar) PDS II monofilament fibers immersed in PBS, pH 7.40, at 50 °C.

Figure 10.

Figure 10

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Comparison of stress at break vs time curves of (blue filled square) uncoated; (yellow filled circle) 0.5, (green filled triangle up solid) 1, (purple filled triangle down solid) 5, and (grey filled tilted square solid) 10 mg/mL LVFX-coated PLC; and (orange filled triangle right-pointing solid) PDS II monofilament fibers immersed in PBS pH 7.40 at 50 °C.

2.9. Assessment of Bacterial Inhibition and Biocompatibility

Since there was no report on the minimum inhibitory concentration (MIC) of LVFX in the CLSI manual for the bacteria in this study; thus, the MIC value was determined. Interestingly, MIC values of LVFX to Pseudomonas aeruginosa, S. aureus (MRSA), and E. coli O157:H7 were approximately similar, while the minimum concentration of LVFX used to inhibit those bacteria was <0.4 μg/mL. Therefore, the final concentration for coating on the suture surface was considered to be between 0.5 and 10 mg/mL, in which the release of the drug after the first 24 h was estimated to be a 10–100 times higher concentration than the MIC.

The LVFX-coated suture inhibited all of the tested pathogenic bacteria at a concentration of 0.5–10 mg/mL, except in S. aureus MRSA wherein only 0.5 mg/mL LVFX was not able to control the bacterial growth (Figures 11 and 12). However, several studies had previously described that S. aureus MRSA resisted various drugs including LVFX; thus, the minimum concentration of this drug released from the suture surface must be at least 2–3 times higher than the MIC. In fact, the inhibition zone was increased in accordance with the greater coating concentration, so to overcome the multidrug-resistant strains, more than 5 mg/mL LVFX was required to coat the suture. In addition, LVFX released from the suture surface reached the MIC value in the first 24 h and the maximum was still detected until day 30 (see Figure 7A,B); thus, the antibacterial shelf-life was determined at only the appropriate coating concentration of LVFX of 10 mg/mL. Interestingly, the inhibition zone for each suture with the single coating was not significantly different from day 1 until day 90 in all tested bacteria (Figure 13). The result suggests that the shelf-life of antibacterial activity could remain for at least 3 months and repeating the coating of LVFX on the surface from 1 to 5 times did not show the significant antibacterial effect on these three pathogenic strains (Table 8). This finding supports the future application of this surface coating production method instead of blending the material together, which would be beneficial in terms of reduction in the overall production cost.

Figure 11.

Figure 11

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Antibacterial activity of the LVFX-coated suture to the representative of three pathogenic bacteria at the concentrations of 0.5, 1, 5, and 10 mg/mL LVFX. The size of the inhibition zone (mm) was investigated at 24 h, and the experiment was performed with independent triplicates.

Figure 12.

Figure 12

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Inhibition zone of LVFX-coated fibers to the representative of three pathogenic bacteria including S. aureus (MRSA) (A1–A3), E. coli O157:H7 (B1–B3), and P. aeruginosa (C1–C3) at the concentration of LVFX = 0.5, 1, 5, and 10 mg/mL. The size of the inhibition zone was investigated at 24 h with independent triplicates using 0.1 mg/mL LVFX and uncoated PLC fiber as the positive and negative controls, respectively.

Figure 13.

Figure 13

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Shelf-life of 10 mg/mL LVFX-coated fibers from day 1 to day 90. The suture was determined for the inhibition zone with the pathogenic bacteria as shown in the picture during the storage time at room temperature.

Table 8. Determination of 10 mg/mL LVFX-Coated Fiber Shelf-Life for Antibacterial Activity for 3 Months with Common Infection Pathogenic Bacteria.

    inhibition zone (mm)
    repeated coating time
tested bacteria time (day) 1 2 3 4 5
S. aureus (MRSA) 1 22.20 ± 0.12 22.60 ± 0.10 20.33 ± 0.58 22.10 ± 0.23 22.00 ± 0.25
  3 22.00 ± 0.23 22.00 ± 0.32 20.33 ± 0.58 23.00 ± 0.00 23.67 ± 0.58
  5 23.33 ± 0.58 23.00 ± 0.00 23.33 ± 0.58 23.00 ± 0.00 23.00 ± 0.00
  7 22.00 ± 0.00 22.67 ± 0.58 22.67 ± 0.58 22.00 ± 0.00 20.00 ± 0.00
  15 20.00 ± 0.00 20.33 ± 0.58 20.33 ± 0.58 21.33 ± 0.58 19.67 ± 0.58
  30 22.00 ± 0.00 23.00 ± 0.00 23.33 ± 0.58 23.33 ± 0.58 24.00 ± 0.00
  60 22.00 ± 0.00 22.33 ± 0.58 22.33 ± 0.58 23.00 ± 0.00 24.00 ± 0.00
  90 21.00 ± 0.00 21.00 ± 0.00 20.67 ± 0.58 22.33 ± 0.58 21.00 ± 0.00
E. coli O157:H7 1 22.23 ± 0.00 22.10 ± 0.20 20.33 ± 0.58 23.00 ± 0.60 23.00 ± 0.98
  3 22.230 ± 0.09 22.80 ± 0.00 20.33 ± 0.58 23.33 ± 0.58 24.00 ± 0.00
  5 22.00 ± 0.00 22.67 ± 0.58 23.33 ± 0.58 23.33 ± 0.58 23.00 ± 0.00
  7 21.00 ± 0.00 21.33 ± 0.58 21.00 ± 0.00 21.67 ± 0.58 21.00 ± 0.00
  15 23.33 ± 0.58 23.00 ± 0.00 21.33 ± 0.58 21.67 ± 0.58 20.67 ± 0.58
  30 21.33 ± 0.58 22.00 ± 0.00 23.00 ± 0.00 23.33 ± 0.58 23.00 ± 0.00
  60 22.33 ± 0.58 22.67 ± 0.58 24.00 ± 0.00 23.67 ± 0.58 24.00 ± 0.00
  90 21.00 ± 0.00 22.00 ± 0.00 23.00 ± 0.00 23.67 ± 0.58 23.33 ± 0.58
P. aeruginosa 1 22.00 ± 0.45 22.00 ± 0.40 20.33 ± 0.58 23.00 ± 0.12 23.40 ± 0.70
  3 22.60 ± 0.45 22.23 ± 0.12 20.33 ± 0.58 23.00 ± 0.00 23.67 ± 0.58
  5 22.00 ± 0.00 22.00 ± 0.00 20.33 ± 0.58 23.33 ± 0.58 23.33 ± 0.58
  7 22.00 ± 0.00 21.33 ± 0.58 21.33 ± 0.58 21.67 ± 0.58 22.33 ± 0.58
  15 21.67 ± 0.58 22.00 ± 0.00 23.33 ± 0.58 23.00 ± 0.00 21.33 ± 0.58
  30 22.00 ± 0.00 23.00 ± 0.00 23.00 ± 0.00 23.67 ± 0.58 23.33 ± 0.58
  60 21.33 ± 0.58 22.00 ± 0.00 21.67 ± 0.58 22.00 ± 0.00 22.00 ± 0.00
  90 23.00 ± 0.00 23.33 ± 0.58 22.00 ± 0.00 22.33 ± 0.58 22.00 ± 0.00
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For investigation of the toxicity of the material and LVFX released from the suture after its incorporation to form the suture and coated suture, the 3-(4,5-dimethyl-2-thiazolyl)-2,5-diphenyl-2H-tetrazoline bromide (MTT) assay was performed with the extracted suture solution. From Figure 14, levofloxacin-coated sutures did not express toxicity to epithelial cell line L929 at all concentrations used in this study since the threshold of viability was greater than 70%. In fact, the maximum LVFX released from the suture on day 30 is from 0.019 to 0.135 mg/mL, which is equivalent to 0.5–10 mg/mL LVFX-coated suture, respectively, did not affect the cell viability. This result confirmed that the suture prototype in this study was not harmful to the cells or probably the tissues.

Figure 14.

Figure 14

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Evaluation of toxicity of coated fibers and LVFX to epithelial cells using the L929 cell line with the MTT assay. The sutures were extracted in PBS solution, and the concentration of LVFX released was mentioned as shown in the figure. The variability above 70% was considered as nontoxic.

3. Conclusions

In this research work and according to ASTM F1925-17 specifications, a medical-grade 70/30 poly(l-lactide-co-ε-caprolactone) (PLC) copolymer was successfully synthesized using 0.01 mol % liquid Sn(OnBu)2 as an initiator at 120 °C for 96 h. It was also found that the copolymer obtained had sufficiently high molecular weight to be melt-spun into monofilament sutures. The mechanical strength of the largely amorphous as-spun fiber could be improved by a controlled sequence of hot-drawing and fixed-annealing steps. Hot-drawing was carried out to induce crystallization and molecular orientation along the fiber axis. Subsequent annealing (at the fixed length) decreased the fiber’s strength (tensile stress at break) but slightly increased its flexibility. The 2nd hot-drawing step after annealing then restored the lost tensile strength. The strongest fiber was obtained from the 2nd hot-drawing process at 90 °C with a stress at break of 434 MPa. In the final step, further thermal treatment by a combination of fixed-annealing and free-annealing stabilized the fiber morphology by allowing molecular relaxation to occur.

For the wound healing enhancement, the fiber was dip-coated with “levofloxacin” by adding the drug into a mixture solution of acetone, poly(ε-caprolactone) (PCL), and calcium stearate (CaSt) in the ratio of acetone/PCL/CaSt = 100:1% w/v:0.1% w/v. The tensile strength of the final coated fiber was found to be increased from about 330 to 400 MPa, which is more or less comparable with that of commercial PDS II of a similar size. Finally, the efficiency of the drug-coated fiber regarding its controlled drug release and antimicrobial activity was investigated, and the results showed that the coated fiber was able to release the drug continuously for as long as 30 days. For fiber antimicrobial activity, it was found that a concentration of 1 mg/mL was sufficient to inhibit the growth of P. aeruginosa, S. aureus (MRSA), and E. coli O157:H7, giving a clear inhibition zone with an average diameter of about 20–24 mm for a duration of more than 3 months. Cytotoxicity testing of the drug-coated fibers showed a nontoxic range of 75–90% viability.

4. Experimental Section

4.1. Materials

l-Lactide was synthesized from l-lactic acid (Jungbunzlauer, 88%) and purified by repeated recrystallization from ethyl acetate and dried to a constant weight in a vacuum oven at 55 °C. The obtained pure l-lactide monomer was a white crystalline solid with a purity ≥99.5% (from DSC purity analysis). Commercial ε-caprolactone (Acros Organics, 99%) was purified by vacuum distillation over calcium hydride. Anhydrous tin(II) chloride (purity > 98%) was purchased from Acros Organics (95%) and used as received. N-Heptane (Carlo Erba Reagents, 99%), diethylamine (Sigma-Aldrich, ≥99.5%), and n-butanol (QRëC, 99.95%) were purified by refluxing with Na metal or CaH2 and distilled before use. Calcium stearate (CaSt, Sigma-Aldrich, 6.6–7.4% Ca basis), acetone (Sigma-Aldrich, ≥99.5%) coating solutions, and levofloxacin antibiotic drug (LVFX, Sigma-Aldrich, ≥99.5%) were used as received. Phosphate-buffered saline (PBS) solution, pH 7.4, was prepared by dissolving a tablet of PBS buffer (Calbiochem, Sigma-Aldrich) in 1 L of deionized H2O. P. aeruginosa, S. aureus (MRSA), and E. coli O157:H7 (American Type Culture Collection (ATCC), Manassa, VA), and Mueller–Hinton agar (MHA, Sigma-Aldrich) were used as received.

4.2. Methods

4.2.1. Synthesis of Liquid Tin(II) n-Butoxide (Sn(OnC4H9)2) Initiator

In this work, the synthesis of liquid tin(II) n-butoxide (Sn(OnC4H9)2) for use as an initiator in the ROP of l-lactide and ε-caprolactone was as described in the modified method by Meepowpan et al.76 It was found that Sn(OnC4H9)2 was obtained as a viscous, dark yellow liquid, in approximately 60–70% yield.

4.2.2. Synthesis of Medical-Grade Poly(l-lactide-co-ε-caprolactone), PLC Copolymer

Medical-grade poly(l-lactide-co-ε-caprolactone), PLC, with a mol % composition of L/C = 70:30 was obtained by ring-opening polymerization (ROP) of l-lactide, LL, in the bulk using 0.01 mol % liquid tin(II) n-butoxide, Sn(OnC4H9)2, as a coordination-insertion initiator. In the experimental procedure, the organotin initiator was added to the reaction flask under dry nitrogen in a controlled atmosphere glovebox at room temperature. After removing the flask from the glovebox, it was immersed in a silicone oil bath at 120 °C for 96 h. At the end of the polymerization period, the flask was removed from the oil bath and allowed to cool down to room temperature. The crude PLC was purified by cutting and grinding into small pieces to increase its surface area and then drying to constant weight in a vacuum oven at 80 °C to remove any trace amounts of residual monomers. According to ASTM F1925-17, the copolymer product was then characterized in terms of its chemical structure, composition, and properties by proton nuclear magnetic resonance (1H NMR) spectroscopy (400 MHz Bruker Avance II NMR spectrometer), dilute-solution viscometry (Schott-Geräte AVS 300, Ubbelohde viscometer no. 532000c), gel permeation chromatography (GPC, Agilent 1260 Infinity II GPC/SEC System), and thermal analysis using differential scanning calorimetry (DSC, PerkinElmer DSC7 Differential Scanning Calorimeter).77 Other properties included specific rotation via polarimetry (ATAGO AP-300 Automatic Polarimeter), residual monomers and solvents by gas chromatography (Agilent Technologies 6850 Series II GC System), residual heavy metals and catalysts by inductively coupled plasma-optical emission spectrometry (Agilent 5110 ICP-OES), and residual water by Karl Fischer titration (Mettler Toledo C30S Compact KF Coulometer).

4.2.3. PLC Fiber Processing by Melt Extrusion

The PLC copolymer was melt-spun into a monofilament fiber using a single-screw extruder with stretching units and associated equipment (Collin Lab & Pilot Solutions GmbH) as shown schematically in Figure 15 using the optimal verified and validated conditions (Table 9). Before melt processing, the PLC resins were dried in a vacuum oven at 80 °C for 24 h. The process started with the polymer granules being fed from a hopper into the barrel of the extruder. The material was then gradually melted at a temperature of 180 °C by the mechanical energy generated by the rotating screw and the heaters arranged along the barrel. The molten polymer was then forced into a single circular die (1 mm diameter), which shaped the molten polymer into a monofilament fiber followed by cooling in a water bath. The as-spun fibers were of uniform diameter in the range of 0.9–1.1 mm and were of clear color with a smooth surface appearance. On their emergence from the die, the as-spun fibers were cooled in a water bath followed by a series of well-controlled off-line hot-drawing and annealing steps. Fiber tensile strength was determined at ambient temperature using a Lloyds LRX+ Universal Testing Machine (Lloyds Instruments Ltd., U.K.) equipped with a 100 N load cell and bollard grips. The initial gauge length of each fiber test specimen was 40 mm, and the test was carried out at a cross-head speed of 40 mm/min.

Figure 15.

Figure 15

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Schematic diagram of a single-screw extruder with stretching units and associated equipment: ① melt pump for Teach-Line E 20 T-MP, ② screw 20 mm diameter × 25 D, ③ die-head 1.0 mm diameter, ④ melt temperature measurement, ⑤ water bath, ⑥ hot water tempering unit, ⑦ hold-back unit MDO-AT, ⑧ hot-air heating oven, and ⑨ stretching unit MDO-BT and winder.

Table 9. Processing Conditions and Parameters Used for Melt Extrusion of PLC Monofilament Fibersa.
conditions/parameters
Temperature Zone (°C)
zone 1—feeding zone (set/act) 40/40
zone 2—barrel zone 2 (set/act) 165/165
zone 3—barrel zone 3 (set/act) 170/170
zone 4—barrel zone 3 (set/act) 180/180
zone 5—melt pump (set/act) 180/180
zone 7—die (set/act) 170/171
die diameter (mm) 1.0
screw speed, n1 (rpm) (set/act) 15/15
melt pump speed, n2 (rpm) (set/act) 5/5
Melt Pressure (bar)
P1—extruder (set/act) 90/90
P2—filter 85
draw ratio MDO-AT: MDO-BT, 1st hot-drawing (draw speed, rpm) 1:5 (5:25)
hot-air oven temp (°C) 60
1st fixed-annealing (temp, time) 80 °C, 5 h
draw ratio MDO-AT: MDO-BT, 2nd hot-drawing (draw speed, rpm) 1:2 (6:12)
(draw speed, rpm) (6:12)
hot-air oven temp (°C) 90
2nd fixed-annealing (temp, time) 80 °C, 24 h
free-annealing (temp, time) 60 °C, 72 h
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a

Note: set/act = set/actual value.

4.2.4. Preparation of Antimicrobial LVFX-Coated PLC Monofilament Sutures

An antimicrobial drug coating solution was prepared by dissolving 1.0 w/v % medical-grade poly(ε-caprolactone) (PCL) and 0.1 w/v % calcium stearate (CaSt) in 100 mL of acetone according to the recommended protocol.78 Subsequently, levofloxacin (LVX) with mass concentrations of 0.1, 0.2, 1.0, and 2.0 wt % was gradually dissolved in the acetone–PCL–CaSt solution and continuously stirred for 1 h to obtain the homogeneous drug-coated solution. For the fiber coating process (Figure 16), first, the PLC fiber was passed through the coating bath that consisted of a series of small roller guides for directing the suture into the coating solution. As the fiber passed from roller guides 1 and 2 to roller guide 3, it is submerged in a coating mixture to provide a wet-coated fiber. Then, the coated suture was pulled through the drying chambers ② and wound on a take-up reel of an MDO-BT ③.

Figure 16.

Figure 16

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Schematic diagram of a dip coating line of PLC fiber: ① stainless steel coating bath with roller guides, ② hot-air heating oven, and ③ take-up unit MDO-BT and winder.

4.2.5. Confirmation of LVFX Coating on PLC Sutures

Scanning electron microscopy (SEM) is an electron optical imaging technique that yields both topographic images and elemental information when used in conjunction with energy-dispersive X-ray fluorescence (EDXRF). Therefore, in this work, SEM (JSM 5910 LV Scanning Electron Microscope, JEOL Ltd., Japan) attached with EDXRF (INCA Energy Software, Oxford Instruments, U.K.) was used to observe the morphology and chemical composition of the coated samples. Gold sputtering was used to make the coating surfaces conductive for the SEM-EDXRF investigations.

To obtain additional structural information, FTIR in an attenuated (ATR) mode was used to confirm the presence of the LVFX coating on the PLC fiber.

4.2.6. Drug Release Behavior of LVFX-Coated PLC Sutures

For the in vitro hydrolytic biodegradation drug release experiments, five sets of uncoated and 0.1, 1.0, 5.0, and 10 mg/mL drug-coated fibers, each about 70 cm length, were accurately weighed and immersed in a screw-top glass bottle containing 10 mL of the phosphate buffer saline (PBS) solution (pH = 7.40 ± 0.2). Then, all bottles were immediately placed in an incubator that was thermostatically controlled at 37.0 ± 1.0 °C. At a certain period of time, the fiber samples were filtered out and dried under vacuum at room temperature to constant weight. After that, drug release behavior of sutures during in vitro degradation was measured.

The adsorption intensity of the LVFX drug was measured by a UV spectrophotometer (PerkinElmer LAMBDA 850+ UV/vis spectrophotometer) at 298 nm using 0.1 M HCl solution as a blank control. The calibration curve of the LVFX drug (see Supporting Information page S3) was obtained from the plots between the drug concentration (mg/mL) and absorbance (AU); therefore, the amount of LVFX in the release medium can be determined by this standard curve and the cumulative release rate was then calculated according to eq 1 as follows79

4.2.6. 1

where Qt is the cumulative release rate of LVFX at time t, and mt and mi are the amount of released LVFX at time t and initial time, respectively.

4.2.7. Accelerated In Vitro Hydrolytic Biodegradation

For the accelerated in vitro hydrolytic biodegradation experiments, five sets of uncoated; 0.1, 1, 5, and 10 mg/mL drug-coated fibers; and commercial long-term monofilament suture “PDS II” (Ethicon), each about 70 cm in length, were accurately weighed and immersed in a screw-top glass bottle containing 10 mL of phosphate-buffered saline (PBS) solution (pH = 7.40 ± 0.2). Then, all bottles were immediately placed in the incubator that is thermostatically controlled at 50.0 ± 2 °C. At a certain period of time, the fiber samples were filtered out and dried under vacuum at room temperature to constant weight. After that, % weight loss and mechanical strength of sutures during the degradation period were measured.80,81

4.2.7.1. Weight Loss Profiles

After the designated time interval for accelerated in vitro biodegradation studies, a bottle containing either uncoated, LVFX-coated, or commercial (PDS II USP 2/0, Ethicon) sutures was taken out from the incubator. The fiber samples were then filtered off, washed with distilled water, and dried to a constant weight in a vacuum oven at room temperature. After that, their weights were accurately recorded and % weight loss can be calculated using eq 2 as follows

4.2.7.1. 2

where wi and wf are the initial and final weights of fiber samples, respectively.

4.2.7.2. Mechanical Strength

Complementary to the weight loss data is the tensile strength profile. Therefore, in this study, the tensile strength of the fiber was conducted in accordance with the ASTM D2256/D2256 M-10(2015).82 Each wet fiber sample was tested at least six times, and the mean values ± standard deviation (SD) were recorded. The % breaking tensile strength retention (% BSR) was calculated as follows

4.2.7.2. 3

where S0 is the strength before degradation and St is the strength after degradation time t.

4.2.8. Assessment of Bacterial Inhibition

In this work, the antibacterial efficacy of LVFX was determined by the defection assay on the solid medium by measurement of the clear zone around the LVFX-coated suture and LVFX disc. For the assessment method, LVFX was determined for its minimum inhibitory concentration (MIC) with three representative pathogenic bacteria including S. aureus, E. coli O157:H7, and methicillin-resistant P. aeruginosa (MRSA). The MIC value of each bacterium was used for the calculation to coat the suture surface from 0.5, 1, 5, and 10 mg/mL for further analysis. LVFX-coated sutures and control were tested in vitro with those aforementioned on Mueller–Hinton agar (MHA). First, individually tested bacterium was cultivated in 5 mL of tryptic soy broth (TSB) at 37 °C for 24 h with shaking (200 rpm). After that, the culture was centrifuged at 7000 rpm for 5 min to collect the cell pellets. Each bacterial cell pellet was suspended in 0.85% NaCl to adjust the turbidity at OD600 to 0.1 or equivalent to 0.5 McFarland standard (approx. 1 × 108 CFU/mL). The culture suspension was separately swabbed on MHA, and the tested samples including 1 cm of LVFX-coated suture, noncoated suture, and 5 μg of LVFX disc were placed on the plates and then incubated at 37 °C for 24 h. For shelf-life determination of the LVFX-coated suture, a 10 mg/mL drug-coated suture was selected to examine the inhibition zone at day 1 to day 90 with the similar method of the direct diffusion assay that described previously. For statistical analysis, the data were compared by t-tests and significance was defined as P < 0.01.

4.2.9. Biocompatibility

For biocompatibility evaluation, LVFX-coated sutures were determined by the MTT assay based on ISO 10993-5 using mouse fibroblast cell line, L929 (ATCC CCL1, NCTC clone 929). The test principle is based on the reduction of (3-(4,5-dimethyl-2-thiazolyl)-2,5-diphenyl-2H-tetrazoline bromide or MTT reagent) to form azure formazan in living cells that enables to evaluate the enzymatic activity of the succinate dehydrogenase to determine the cellular vitality. For cytotoxicity evaluation, both uncoated and coated sutures at various concentrations of 0.1, 1.0, 5.0, and 10 mg/mL were examined. In addition, LVFX at the concentrations of 0.019, 0.036, 0.070, and 0.135 mg/mL that are equivalent to the maximum concentration released from suture on day 30 from each LVFX-coated suture were used as the control. L929 cell line suspension (1 × 105 cells/mL) in MEM completed medium was seeded into the 96-well plate and was incubated at 37 °C, 5% CO2, and 95% relative humidity for 24 h to obtain confluent monolayers of cells prior to testing. The MEM completed medium was replaced with the extracts of tested sterile uncoated, coated suture extracts, and LVFX solution. For extraction, 0.1 g of sutures were extracted by dissolving in 1 mL of sterile phosphate-buffered saline (PBS) solution at 37 °C for 24 h. Then, 10 μL of the extraction solution or LVFX solution was added to the cell suspension and incubated as the similar protocol mentioned previously. The polyurethane film containing 0.1% zinc diethyldithiocarbamate (ZDEC) was used as a positive control, while the medium without specimen was set as a blank. After 48 h, 10 μL of MTT (5 mg/mL) solution was added and further incubated for 2 h. After removing the MTT solution, 50 μL of dimethyl sulfoxide (DMSO) was added to each well and incubated for 10 min, and then, the absorbance of the suspension was measured at 570 nm. All experiments were performed in triplicate, and the relative cell viability (%) was expressed as a percentage relative to the untreated control cells (blank).

Acknowledgments

This work was financially supported by the Program Management Unit for Enhancement of the Country’s Competitiveness (PMU-C); Office of National Higher Education, Science Research and Innovation Policy Council (NXPO), Thailand. The authors also thank the CMU Presidential Scholarship 2021 for partially funding this research project.

Supporting Information Available

The Supporting Information is available free of charge at https://pubs.acs.org/doi/10.1021/acsomega.1c03569.

  • Results and Discussion—Figure S1, 400 MHz 1H NMR spectrum of a medical-grade PLC synthesized using 0.01 mol % liquid Sn(OnC4H9)2 as an initiator at 120 °C for 96 h; Figure S2, GPC chromatogram of the synthesized PLC copolymer using THF as a solvent at 40 °C and calibrated with polystyrene standards; Figure S3, FTIR spectrum in the region between 4000 and 400 cm–1 of a PLC fiber; and Figure S4, FTIR spectrum in the region between 4000 and 400 cm–1 of a 10 mg/mL LVFX-coated PLC fiber; Experimental Section—Table S1, standard concentration curve for LVFX-coated PLC fiber; and Figure S5, standard concentration plots between LVFX-coated PLC fiber with various concentrations and absorbance at 298 nm (PDF)

Author Contributions

All authors made significant contributions to this work.

The authors declare no competing financial interest.

Supplementary Material

ao1c03569_si_001.pdf (429.1KB, pdf)

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