Skip to main content
NCBI home page
ACS AuthorChoice logoLink to ACS AuthorChoice
. 2024 Feb 20;16(8):11062–11075. doi: 10.1021/acsami.3c18093

Self-Reinforced PTLG Copolymer with Shish Kebab Structures and a Bionic Surface as Bioimplant Materials for Tissue Engineering Applications

Jiafeng Li †,§,∥, Pengfei Jiang †,§,∥, Jianwei Yang †,§,∥, Quntao Zhang †,§,∥, Huiyuan Chen †,§,∥, Ziyue Wang †,§,∥, Chang Liu †,§,∥, Tiantang Fan #,*, Lu Cao ¶,⊥,*, Junhui Sui ‡,*
PMCID: PMC10910444  PMID: 38378449

Abstract

graphic file with name am3c18093_0015.jpg

Green and biodegradable materials with great mechanical properties and biocompatibility will offer new opportunities for next-generation high-performance biological materials. Herein, the novel oriented shish kebab crystals of a novel poly(trimethylene carbonate-lactide-glycolide) (PTLG) vascular stent are first reported to be successfully fabricated through a feasible solid-state drawing process to simultaneously enhance the mechanical performance and biocompatibility. The crystal structure of this self-reinforced vascular stent was transformed from spherulites to a shish kebab crystal, which indicates the mechanical interlocking effect and prevents the lamellae from slipping with a significant improvement of mechanical strength to 333 MPa. Meanwhile, it is different from typical biomedical polymers with smooth surface structures, and the as-obtained PTLG vascular stent exhibits a bionic surface morphology with a parallel micro groove and ridge structure. These ridges and grooves were attributed to the reorganization of cytoskeleton fiber bundles following the direction of blood flow shear stress. The structure and parameters of these morphologies were highly similar to the inner surface of blood vessels of the human, which facilitates cell adhesion growth to improve its proliferation, differentiation, and activity on the surface of PTLG.

Keywords: poly(trimethylene carbonate-lactide-glycolide); solid-state drawing; orientation, stress-induced crystallization; mechanical performance; biocompatibility

1. Introduction

Poly(lactic acid) (PLA) with great mechanical properties and biocompatibility offers new opportunities for next-generation high-performance biological materials.1−6 At present, PLA is used in fully biodegradable medical devices (such as orthopedic internal fixation, vascular stents, etc.), drug delivery systems, and tissue engineering.7−9 As is known, the ideal biodegradable polymer materials should have excellent mechanical properties, biocompatibility, a suitable degradation rate, nontoxic degradation products, and good processing and molding performance. However, PLA exhibits unsatisfactory performances such as insufficient mechanical strength and biocompatibility.10−13 Therefore, the research for high-performance and functional PLA materials not only provides a theoretical basis for the application of new biomedical material design, structure, and performance control mechanism but also provides the premise of developing PLA biomedical devices with excellent comprehensive performance. In recent years, various technical routes for simultaneously improving both mechanical strength and biocompatibility have been attempted for PLA.

The modification of PLA mainly involved the modification of the bulk and surface. The main modification of PLA was to modify the crystallization performance, mechanical properties, melt strength, heat resistance, degradation performance, etc. The surface modification of PLA mainly changed the surface hydrophobicity, surface morphology, surface groups, and bioactive factors. Improving the mechanical properties of polymer materials through chain orientation was an effective and economical method for manufacturing self-reinforcing polymer materials, including blow molding, biaxial stretching, melt spinning, and extrusion. Yu et al. prepared the oriented PLA films with a draw ratio of 200% by the film blowing process.14 Gupta et al. reported the preparation of PLA fibers with a tensile strength of 160 MPa by the melt spinning method under the condition of a draw ratio of 300%.15 The solid-state drawing (SSD) technology can not only significantly improve the mechanical strength but also introduce a special morphology on the material surface, which facilitates the cell adsorption on the material surface.

The SSD technology can not only significantly improve the mechanical strength as the conventional processing methods do but also introduce a special morphology on the surfaces of the materials, which is believed to facilitate cell adsorption on the surfaces of the materials. However, the method of self-enhancement could improve the strength but not the toughness.

At present, there are chemical modification and physical modification methods for toughness modification, mainly including copolymerization modification, grafting modification, and blending modification.16−20 Among them, copolymerization modification is the most common technical means to improve the comprehensive performance of PLA materials. Poly(1,3-trimethylene carbonate) (PTMC) is a common elastomer with good biocompatibility and cell permeability and is widely used in the biomedical field. The glass transition temperature of PTMC is −20 °C, indicating excellent flexibility. Meanwhile, in living organisms, PTMC can be rapidly degraded by biological enzymes into neutral degradation products such as 1,3-propanediol and CO2, which can neutralize the acidic products of PLA during the degradation process and reduce the occurrence of inflammatory reactions. Fan T. et al. synthesized poly(l-lactic acid (PLLA)–PTMC binary copolymers and pointed out that the introduction of PTMC can effectively improve the toughness of PLLA materials.21,22 Besides, the degradation rate of biomedical materials is also a very important factor in evaluating the performance of medical device materials. According to the preliminary work results of this project, it was found that the (glycolide) GA unit has high reactivity and is prone to the polymerization reaction with the LLA unit. The addition of the GA unit can greatly disrupt the chain regularity of PLLA, reduce its crystallinity, and accelerate the degradation rate of PLLA copolymer materials.23,24

In this paper, using the SSD process, we demonstrate a novel poly(trimethylene carbonate-lactide-glycolide) (PTLG) material with oriented shish kebab crystals and bionic surface structures based on copolymerization of LLA, TMC, and GA. The toughness was improved by introducing flexible TMA units into the LA molecular chain, and the introduction of GA units increased the degradation rate. On one hand, the crystal structure of this self-reinforced vascular stent was transformed from spherulites to shish kebab crystallite. The shish kebab structure can prevent the lamellae from slipping and produce a mechanical interlocking effect, which could significantly improve the mechanical strength of the materials. On the other hand, the stress-induced crystallization effect formed fibrous crystals or shish kebab crystals, which exhibit a micro ridges and grooves morphology on the material surface. The structure and parameters of these morphologies were highly similar to the inner surface of blood vessels of the human. These ridges and grooves were attributed to the reorganization of cytoskeleton–fiber bundles following the direction of blood flow shear stress. This special surface topography was similar to the surface of the inner wall of the blood vessel, resulting in bionic characteristics. The bionic surface significantly promoted the adsorption growth and proliferation of cells on the surface of the material. The SSD process of PTLG samples provides ideas for the research and preparation of novel vascular stents with excellent comprehensive properties. It provides a novel processing technology for the biomaterials of implant devices Figure 1.

Figure 1.

Figure 1

Open in a new tab

Synthetic route of the PTLG random copolymer.

2. Materials and Methods

2.1. Preparation of The PTLG Copolymers

PTLG copolymers were synthesized by ring opening polymerization. l-Lactide (LLA) and glycolide (GA) monomers were heated and dissolved in ethyl acetate and then naturally cooled and recrystallized to remove impurities and moisture. Stannous octanoate [Sn (Oct)2] was distilled under reduced pressure at 140 °C for 90 min before use to remove residual moisture from the system. Certain amounts of LLA, GA, and TMC were put into a polymerization tube using Sn(Oct)2 as a catalyst, the mass ratio of the monomer to catalyst was 3000:1, and the remaining catalyst in the prepared PTLG was measured by inductively coupled plasma mass spectrometry, which indicated that the remaining catalyst in the prepared PTLG was 5 ppm. The feed ratios of TMC, LLA, and GA were 0:95:5, 1:94:5, 3:92:5, and 5:90:5. The polymerization reaction was done at 135 °C for 72 h. The initial product obtained was dissolved in CHCl2, and it was purified with ethanol as a precipitant to remove unreacted monomers. The purified samples were dried under vacuum at 60 °C to a constant weight to obtain a white solid sample. All the copolymers were obtained with yields higher than 85%. PTLG5/90/5 indicates that the molar ratio of TMC, LLA, and GA in the PTLG random copolymer was 5:90:5. The unit ratios of the obtained PTLG and the monomer ratio in feed are shown in Table 1.

Table 1. Characterization of PTLG Copolymers.

samples feed ratio product molar ratio M̅n × 10–4 (g/mol) M̅w × 10–4 (g/mol) PDI
PTLG5/95/0 5/95/0 5.6/94.4/0 8.45 9.65 1.145
PTLG5/94/1 5/94/1 4.5/94.3/1.2 8.45 10.41 1.232
PTLG5/92/3 5/92/3 4.9/91.5/3.6 7.51 8.87 1.181
PTLG5/90/5 5/90/5 5.8/90.6/4.2 10.40 15.19 1.460
Open in a new tab

2.2. Solid-State Drawing of PTLG Copolymers

After the sample was synthesized, it was dissolved and poured into a film, chopped and dried, and then added to an extruder to extrude and pull into the wire. The wire was oriented through SSD, and then, the oriented wire was woven into a vascular stent. First, wire samples were prepared by the extrusion–traction–filament method. Second, SSD was performed by fixing the samples on a fixture in a constant temperature box, preheating for 30 min, starting the motor, and conducting the SSD process. The drawing temperature was 120 °C and the stretching rate was 60 mm/min. After the samples reached the design length, the drawing was stopped, cooled, and unloaded to obtain the SSD wire. Finally, the SSD wire was woven to prepare a vascular stent.

2.3. Characterization of PTLG Copolymers

2.3.1. Two-Dimensional Small-Angle X-Ray Scattering

Two-dimensional small-angle X-ray scattering (2D SAXS) was carried out at the BL16B of the Shanghai Synchrotron Radiation Facility (SSRF) with a wavelength of 0.124 nm and a detector-to-sample distance of 1840 mm.

2.3.2. Two-Dimensional Wide-Angle X-Ray Diffraction

Two-dimensional wide-angle X-ray diffraction (2D WAXD) was carried out at the BL16B of the SSRF with a wavelength of 0.154 nm and a detector-to-sample distance of 154 mm with Cu Kα radiation.

2.3.3. Mechanical Properties

Tensile testing was carried out at 5 mm/min on a universal testing machine (Instron 5943) at room temperature.

2.3.4. Scanning Electron Microscopy Analysis

The morphologies of the samples were characterized with a Nanoscope IV scanning electron microscope (Veeco), and the voltage was 15 kV.

2.3.5. Cell Culture

L929 cells were purchased from the Shanghai Cell Bank of the Chinese Academy of Sciences. The details of the cell culture are shown in the Supporting information.

2.3.6. Cell Proliferation

The cell proliferation of the SC-PLA samples was assessed by the cell counting kit-8 (CCK-8) assays, and the details are shown in the Supporting information.

2.3.7. Cell Live/Dead Viability Assay

Calcein-AM and propidium iodide (PI) were used to perform the cell Live/Dead viability assay, and the details are shown in the Supporting information.

2.3.8. Cell Nucleus/Cytoskeleton Staining

Cell nucleus/cytoskeleton staining was performed by using the phalloidin conjugate solution, and the details are shown in the Supporting information.

2.3.9. Statistical Analysis

Statistical analysis and calculations were performed according to SPSS.

3. Results and Discussion

The self-reinforced modification method of SSD had important research value for the application of biomaterials in the biomedical field. The SSD process can enhance the orientation of rods, pipes, wires, and sheets and thus prepare high-performance biomedical products through turning, laser engraving, weaving, and other methods, such as orthopedic internal fixation products (bone screws, bone plates, and bone needles), vascular stents, surgical sutures, etc.

There were two main ways to prepare self-reinforced PTLG vascular stents using SSD. For pipes, the processing method included orientation processing followed by laser engraving. PTLG samples were extruded into a tube by an extruder and then processed by SSD to prepare a highly oriented tube, which was then mechanically cut or laser engraved to prepare the vascular stent. For wires, the processing method included orientation processing followed by weaving. PTLG samples were extruded into a wire by extrusion and traction and then processed by SSD to prepare a highly oriented wire, which was then subjected to weaving to prepare the vascular stent. Our study used the method of wire orientation processing followed by weaving to prepare vascular stents.

The schematic of the self-reinforced PTLG vascular stents with shish kebab structures and a bionic surface s shown in Figure 2. The PLTG random copolymer was prepared by ring-opening polymerization. Its toughness was improved by introducing flexible TMA units into the LA molecular chain, and the introduction of GA units increased the degradation rate. Furthermore, the oriented PTLG monofilament was prepared by the SSD process, and the monofilament was woven into a self-reinforced PTLG vascular stent. On one hand, the crystal structure of this self-reinforced vascular stent was transformed from spherulites to shish kebab crystallite. When the materials were damaged by external force, the staggered arrangement between the adjacent shish kebab structure can prevent the lamellae from slipping and produce a mechanical interlocking effect, which could significantly improve the mechanical strength of the materials. On the other hand, during the SSD process, the stress-induced crystallization effect caused the initial spherical crystals of PTLG samples to slip, rupture, and recombine to form fibrous crystals or shish kebab crystals, which exhibit a micro ridges and grooves morphology on the material surface. The structure and parameters of these morphologies were highly similar to the inner surface of blood vessels of the human. The surface topography of the self-reinforced vascular stent changed from flat to a parallelly arranged microgroove structure. These ridges and grooves were attributed to the reorganization of cytoskeleton–fiber bundles following the direction of blood flow shear stress. This special surface topography was similar to the surface topology of the inner wall of the blood vessel, resulting in bionic characteristics. The bionic surface significantly promoted the adsorption growth and proliferation of cells on the surface of the material. Therefore, the toughness, degradation performance, mechanical strength, and biocompatibility of the self-reinforced vascular stent have been improved by random copolymerization and orientation processing, which will provide ideas for the research and preparation of new vascular stents with excellent comprehensive properties.

Figure 2.

Figure 2

Open in a new tab

Schematic of the self-reinforced PTLG vascular stents with a shish kebab structure and a bionic surface.

3.1. Chemical Structure of The PTLG Materials

The important factor affecting medical polyester materials was molecular weight (M̅n). When M̅n was lower than 2.5 × 104 g/mol, the mechanical strength of the materials will not be able to meet the patient’s needs. On the other hand, when the clinical polyester materials were processed by extrusion and injection, it inevitably led to heat degradation due to the small amount of water and the shear stress of the screw, such that the M̅n of materials decreased. Taking into account the above factors, this experiment prepared a higher-molecular-weight PTLG random copolymer by controlling the experimental conditions and standardized operations, and its gel permeation chromatography (GPC) curve is shown in Figure 3a. M̅n, M̅w, and the molecular weight distribution index (PDI) of PTLG copolymers are shown in Table 1. It can be seen that the M̅n of the PTLG random copolymer was about 7.45–10.40 × 104 g/mol, the M̅w of the PTLG random copolymer was about 8.87–15.19 × 104 g/mol, and the molecular weight distribution index was between 1.145 and 1.460. Among the PTLG samples, PTLG/5/90/5 exhibited the highest molecular weight of 10.40 × 104 g/mol, which promoted the mechanical strength and processing performance. Therefore, PTLG/5/90/5 samples were selected for further orientation processing by SSD.

Figure 3.

Figure 3

Open in a new tab

(a) GPC curves of PTLG copolymers; (b) FTIR spectra of PTLG copolymers; (c) 1H NMR spectra of PTLG copolymers; and (d) 13C NMR spectra of PTLG copolymers.

The Fourier transform infrared (FTIR) spectra of the PTLG random copolymer are shown in Figure 3b. From the FTIR spectrum of PTLG5/90/5, it can be seen that the characteristic peak at 1077 cm–1 belongs to the stretching vibration peak of –CH3 in PLLA, while the characteristic peak at 867 cm–1 belongs to the stretching vibration peak of–C–C in PLLA. The stretching vibration peak of C–O–C in PTMC is split into two characteristic peaks: 1179 and 1264 cm–1, corresponding to the characteristic absorption peaks of PLLA and PTMC, respectively, indicating that the LLA monomer is chemically bonded with TMC. Meanwhile, the characteristic peak at 1449 cm–1 corresponds to the bending vibration peak of –CH2 in the GA unit. The FTIR spectroscopy results also indicate the synthesis of PTLG random copolymers.

1H NMR can be used to analyze the chemical structure. The 1H NMR spectra of the PTLG random copolymer are shown in Figure 3c, and it can be seen that the PTMC spectrum has a peak located at peak 2 (δ = 4.23 ppm) and a peak at peak 3 (δ = 2.02 ppm), which belonged to the H atoms of –CH2 at both ends of PTMC and the H atom of –CH2 in the middle, respectively, and the intensity of the characteristic peaks is relatively low. This is mainly due to the low content of TMC, which is consistent with the FTIR spectroscopy results of PTLG random copolymers. Peak 1 (δ = 5.15 ppm) and peak 4 (δ = 1.7 ppm) were attributed to the H atoms in –CH and –CH3 on the PLLA chain segment, respectively. When the GA unit content is 0 mol %, the peaks of PTLG0/95/5 were observed at δ = 5.0–4.5 ppm. In the magnified spectrum of δ = 5.0–4.5 ppm, complex multiple peaks can be observed in the PTLG copolymer when GA units are introduced. According to the analysis of 1H NMR of PTLG block copolymers, the multiple peaks belong to the characteristic peaks of the H atom of –CH2 in the GA unit. The main reason for the appearance of multiple peaks is due to the high reactivity of the GA unit, which makes it easy for the LLA and TMC units to undergo polymerization reactions, resulting in a decrease in the regularity of the PLLA chain segments.

The proportion of each component in the PTLG random copolymer was calculated based on the characteristic peaks of the H atom in –CH2 at both ends of TMC, the characteristic peaks of the H atom in –CH on LLA unit, and the integrated intensity of the H atom in –CH2 on the GA unit. The results are listed in Table 1, and it can also be seen that the actual composition ratio of each component of the PTLG random copolymer was basically consistent with the original feed ratio.

The 13C NMR spectra of PTLG copolymers are shown in Figure 3d, and it can be seen that the chemical shifts of the C atom in the C=O of TMC, LLA, and GA units were 154.4 ppm (peak 2), 169.8 ppm (peak 1), and 167.1 ppm (peak 7), respectively. The peak 5 of δ = 27.7 ppm represents the chemical shift of C atoms in the middle –CH2 and the –CH2 at both ends of the TMC unit, respectively. The chemical shift of the C atom in the PLLA main chain –CH was located at 68.9 ppm (peak 3), and it was split into three weaker characteristic peaks, δ = 69.3, 69.2, and 69.1 ppm, respectively. This was mainly due to the introduction of GA units disrupting the regularity of PLLA chain segments. The chemical shift of the C atom in the PLLA side chain –CH3 was located at 16.8 ppm (peak 6). The C atom in GA unit –CH2 was located at δ = 60.9 ppm (peak 4) and was clearly split into multiple peaks. From the above analysis, it can be seen that the C atom in both LLA and GA units exhibited multiple characteristic peaks, directly indicating that the LLA and GA units in PTLG random copolymers existed through chemical bonding and have multiple sequence structural units.

Based on the previous analysis results, the relationship between the average sequence length and performance of different units in copolymers was analyzed for 60–72 ppm, and the results are shown in Figure 4 and Table 2. Taking PTLG5/90/5 as an example, the average sequence length of each unit in the PTLG block copolymer was calculated based on the integrated intensity of the characteristic peaks in Figure 4. The calculation formulae are as follows

3.1. 1
3.1. 2
3.1. 3

Figure 4.

Figure 4

Open in a new tab

13C NMR spectra of PTLG5/90/5 (partial enlargement).

Table 2. Assignment of Sequences in the 13C NMR Spectra of the PTLG Copolymera.

number sequence chemical shift (ppm)
Lactide Sequences
1 TLLT 71.33
2 TLLG 69.26
3 TLLLL + TLLLT 69.21
4 TLLG 69.16
5 LLLL 69.01
6 LLGG 68.85
Carbonate Sequences
7 GT’G + TT’G 64.97
8 TT’T + TTT’ 64.26
9 TT’GG 64.20
11 GGT’’T 61.87
Glycolide Sequences (Methylene Carbon Region)
10 TGT + TGGT 63.50
12 GGLL 60.88
13 GGGG 60.80
14 LGL 66.67
Open in a new tab
a

T′ = –OCH2CH2CH2OCO– and T′′ = –OCH2CH2CH2OCO–.

Here, lLL, lT, and lGG are the average sequence lengths of LLA, TMC, and GA units, respectively. They represent the integrated intensity of each characteristic diffraction peak.

The relationship between the average sequence length and performance of different units in PTLG copolymers was studied. The average sequence length of each unit in the PTLG random copolymer is listed in Table 3. It can be seen that the GA content had a significant impact on the average sequence length of LLA. As shown in Table 3, when the GA content was 0 mol %, the lLL value in PTLG0/95/5 was 18.9. When the GA content increases, the lLL value in the PTLG copolymer gradually decreased, such as the lLL value in PTLG1/94/5 decreasing to 11.7, while when the GA content increases to 5 mol %, the lLL value in PTLG5/90/5 decreased to 3.9. Meanwhile, the average sequence length of TMC units gradually decreased, but the amplitude was relatively small. The average sequence length of GA units remained almost unchanged. This was mainly due to the higher reactivity of the GA unit than that of the LLA and TMC units.

Table 3. Microblock Lengths of the Long Block for PTLG Copolymers.

samples lLLb lTb lGGb
PTLG0/95/5 18.6 2.5 0
PTLG1/94/5 11.7 2.3 0.6
PTLG3/92/5 5.7 2.1 0.7
PTLG5/90/5 3.8 2.0 0.9
Open in a new tab

3.2. Microstructure Characteristics of the PTLG Materials by the SSD Process

The effect of different draw ratios on the orientation structure of PTLG random copolymers was studied using a solid-state stretching device. The introduction of the TMC unit structure enhanced the flexibility of the chain and endowed PTLG with high entanglement density and melt strength. The draw ratio of the PTLG random copolymer was as high as 1200%, higher than the maximum value of 600% in linear PLA experiments.25,26 The 2D WAXD patterns of samples with different draw ratios during SSD are shown in Figure 5. From Figure 5, it can be seen that due to the random distribution of spherulites in the samples, it was proven that the initial PTLG random copolymer did not exhibit the pattern of Debye–Scherrer diffraction rings. However, through SSD processes with different draw ratio parameters, two different diffraction points in the equatorial direction of the sample (200)/(110) reflection ring and four concentrated arcs in the reflection ring (203) were detected, indicating that orientation behavior occurred within the microstructures of polymer chains and crystals. As the draw ratio increased from 600 to 1200%, the orientation structure developed. These circular diffraction rings gradually became concentrated arcs and ultimately exhibited different diffraction points.

Figure 5.

Figure 5

Open in a new tab

2D WAXD patterns of the PTLG samples with different draw ratios.

In order to obtain more intuitive information about the crystal structure of the samples, the 2D WAXD patterns were converted to the general WAXD curves by using 2θ integration, as shown in Figure 6a. The diffraction peak at 16.7° and α′ disordered crystal structure was related, corresponding to the (200)/(110) crystal planes, respectively. For the PTLG-100% samples, the diffraction peaks of the (200)/(110) crystal planes were weak, indicating the lower crystallinity of the samples. When the draw ratio exceeded 500%, it was observed that the diffraction peak intensity of the samples gradually increased, indicating the presence of stress-induced crystallization during the SSD process, which facilitated the crystallization and increased the crystallinity of the PTLG samples.

Figure 6.

Figure 6

Open in a new tab

(a) One-dimensional WAXD curves of the 2D WAXD patterns; (b) azimuthal intensity distributions of 2D WAXD patterns; (c) DSC curves of the PTLG samples; and (d) crystallinities and orientation degrees of the PTLG samples.

As shown in Figure 6b, the 2D WAXD patterns were integrated as a function of the azimuth angle between 0 and 360°. For PTLG random copolymer samples prepared by the SSD process, significant peaks were observed at azimuthal angles of 90 and 270°, indicating a high orientation degree. It can be seen that the PTLG-100% samples exhibited a weak azimuthal integration peak, indicating a low orientation degree and random crystal arrangement. As the drawing ratio increased, the intensity of the azimuthal integral peaks located at 180 and 360° gradually increased, indicating that the lamellae were oriented and regularly arranged along the drawing direction.

The crystallization behavior of samples with different draw ratios was investigated by differential scanning calorimetry (DSC), as shown in Figure 6c. The initial PTLG-100% samples showed cold crystallization peaks at about 60–80 °C, which are disappeared for the PTLG random copolymer with the draw ratio increased by the SSD. Meanwhile, all of the PTLG random copolymer samples by the SSD process exhibited an endothermic peak. With the draw ratio value of the samples increasing, the intensity of the melting peak gradually increased. The results proved the stress-induced crystallization behavior of samples. Meanwhile, the melting peak intensifies and the top shifts to the lower temperature side, which may be due to the fact that the newly formed crystals during SSD were molecular chains arranged in a regular manner under stress induction, and the crystallization temperature was lower than that of molten crystallization. The lower crystallization temperature resulted in incomplete arrangement of molecular chains, and the thickness of the lamellae formed by stress-induced crystallization was lower than that of molten crystallization. Therefore, the melting peak intensifies and the top shifts to the lower temperature side.

The orientation factor and crystallinity are summarized in Figure 6d. It can be seen that the initial orientation and crystallinity of the samples were 0 and 6.8%, respectively, indicating that the molecular chains of the initial samples exhibited an isotropic random arrangement and a relatively small proportion of the crystal structure. As the draw ratio increased, the orientation degree and crystallinity of the samples gradually increased, indicating that stress-induced crystallization occurred during SSD, where the molecular chains were arranged in an orderly manner to form lamellae, and the maximum orientation factor and crystallinity achieved were 0.9311 and 53.6%, respectively. The increasing of the orientation degree and crystallinity facilitated the enhancement of the mechanical strength of the PTLG samples.

Further characterization of the crystal structure of the samples was carried out using 2D SAXS, as shown in Figure 7. The PTLG-100% samples exhibited isotropic scattering rings, indicating the disordered arrangement of lamellas. When the draw ratio reached 300%, it can be clearly observed that the uniform scattering ring of the samples evolved into the maximum spindle pattern along the meridian direction, which belonged to the shish structure. This indicated that the crystal orientation was arranged and had strong periodicity. When the draw ratio exceeded 700%, not only the maximum spindle pattern along the meridian direction of the samples but also the diffraction bright spots were observed on both sides of the meridian, which belonged to the kebab structure, indicating the formation of a shish kebab crystal structure. In the SSD process, the orientation of the high molecular weight chains formed the shish structure. Then, as the SSD continued, the already formed shish structure was used as the crystal nucleus, and the kebab structure formed and grew on the shish structure on both sides. Therefore, the bright spots on both sides of the meridian appeared later than the maximum spindle pattern in the meridian direction.

Figure 7.

Figure 7

Open in a new tab

2D SAXS patterns of the PTLG samples with different draw ratios.

The one-dimensional scattering intensity distribution curves of the samples were obtained by integrating the 2D SAXS pattern, as shown in Figure 8a. The peak position of the curves was calculated based on the formula Lac = 2π/qmax to determine the long period (Lac) of the samples. It can be seen that the peak position of the one-dimensional scattering intensity distribution curve of the samples shifted slightly to the right with the increase of the draw ratio, indicating that the long period of the PTLG samples decreases after SSD. Moreover, the scattering intensity distribution correlation function K(z) was obtained from the one-dimensional scattering intensity distribution curves as follows27,28

3.2. 4

where z is the drawing direction. The K(z) curves of PLTG with different draw ratios are shown in Figure 8b, and the positions of the first peak minimum and maximum values, respectively, correspond to the lamellar thickness (Lc) and the long period (Lac). As the draw ratio increased, the maximum position of the peak gradually moved to the left, suggesting the long period decrease. Meanwhile, the minimum position of the K(z) curves remained at a constant value, suggesting that the lamellar thickness remained constant. According to Lac = Lc, + La, the calculated Lac, Lc, and La values are shown in Figure 8d. With the draw ratio increasing, Lac gradually decreases and Lc remains basically constant, proving stress-induced crystallization. According to the results of 2D WAXD and 2D SAXS, in the SSD process, the orientation of the high molecular weight chains formed the shish structure, and then, the kebab structure formed and grew on the shish structure on both sides, forming the shish kebab structure. The changes of Lac and Lc were consistent with the DSC results.

Figure 8.

Figure 8

Open in a new tab

(a) One-dimensional scattering intensity distribution curves of PTLG samples; (b) correlation function K(z) curves; (c) K(z) curves of PTLG samples; and (d) the long period (Lac), lamellar thickness (Lc), and amorphous region thickness (La) of PTLG samples.

The initial PTLG samples presented a typical spherical morphology, with clear boundaries between the spherulites. The size of the spherulites was about 5–10 μm. There was a certain degree of radial growth compression between adjacent spherulites, and the radial growth of crystal clusters was observed in the spherulites. The stacking of crystal clusters was regular and dense. The crystal structure of the PTLG samples subjected to SSD was completely altered, with no spherical crystals present and exhibiting a parallel and relatively regular shish kebab crystal structure. In addition, it can be seen that fibrous crystals along the stretching direction act as crystal nuclei, with regularly arranged crystal structures growing on both sides. The diameter of the fibrous crystal nucleus is 1.5 μm, and the length exceeds 30 μm. The shish kebab crystal structure caused the interlaced arrangement of various crystal chains inside the polymer crystal to form mechanical interlocking, preventing crystal slip and producing a self-enhancement effect, thereby significantly improving the mechanical strength of the PTLG samples.

After the amorphous phase was etched, the shish kebab structure of stress-induced crystallization during the SSD process was examined by scanning electron microscopy (SEM), as shown in Figure S1. The crystal morphology of the initial PTLG samples was spherical, and the size was about 5–10 μm, which indicated that topical shish kebabs were formed for PTLG samples by the SSD process, as shown in Figure S1. It can be seen that the fibrous crystals of the shish along the draw direction acted as the crystal nuclei, and many kebabs with a regularly aligned lamellar structure grew on both sides of the central shish. The self-reinforcing effects of the shish kebab superstructure through a mechanical interlocking of the kebabs inside the semicrystalline polymer was considered to significantly prevent lamellar slippage and cracking.

3.3. Mechanical Properties of the PTLG Materials by the SSD

The stress–strain curves of tensile tests and the mechanical properties of the PTLG samples are shown in Figure 9, respectively. The specific data and error range on mechanical properties for PTLG materials are shown in Table S1. The tensile strength, modulus, and elongation at break of the initial PTLG random copolymer samples were 45.4 MPa, 0.48 GPa, and 34.1%, respectively. With the draw ratio increasing, the elongation at break significantly increased and stabilized at 22.4%. When the draw ratio increased to 1200%, the tensile strength and modulus gradually increased, reaching 333 MPa and 1.93 GPa, respectively. The stress–strain curves of polymer materials mainly exhibited five types: hard yet brittle, hard yet tough, hard yet strong, soft yet tough, and soft yet weak. As a semicrystalline polymer, the stress–strain curve of PTLG mainly presented the characteristics of hardness and toughness. Its stress–strain curve mainly showed five stages: elastic deformation, yield, development of large deformation, strain hardening, and fracture. In the early stage of deformation, the samples mainly showed general elastic deformation. After the samples reached the yield point, the original crystalline structure in the crystalline polymer was destroyed, followed by thin neck yielding and molecular chains and segments undergoing orientation. Crystals underwent slip, fracture, recombination, and recrystallization. Continuing to increase strain promoted the recombination of deformation concentration areas into new, well oriented, and high-strength crystalline structures. As this new structure increases, the stress increases again until it fractures. Therefore, the stress increased beyond the yield point. It was attributed to stress-induced crystallization. The trend of mechanical properties was also in good agreement with the structural evolution.

Figure 9.

Figure 9

Open in a new tab

Stress–strain curves of the PTLG samples with different draw ratios.

Radial support force referred to the resistance of a vascular stent to radial pressure, which determined whether the stent can adhere to the vascular wall after implantation. It was one of the important factors for evaluating the comprehensive performance of vascular stents. Figure 10 shows the pressure displacement relationship curves of thePTLG peripheral vascular stent during balloon dilation experiments. From Figure 10, it can be seen that the PTLG peripheral vascular stent was divided into three stages during balloon dilation. The first stage referred to the slow expansion stage of the stent during the balloon expansion experiment, which was referred to as the elastic deformation stage of the stent during the initial expansion stage; the second stage referred to the rapid expansion process of the PTLG peripheral vascular stent during the midexpansion stage, which was referred to as the plastic deformation stage of the stent during the midexpansion stage; the third stage referred to the rebound stage of the stent when unloading the balloon, which was understood as the morphological change of the stent material due to plastic deformation in the second stage during the balloon expansion experiment. The radial support force of the vascular stent showed an increasing trend with the increase of the drawing ratio, where the radial support force of the PTLG peripheral vascular stent material increases from 12.6 to 35.9 N, and the trend of radial force was also in good agreement with the structural evolution. The increase of the radial support force promoted the vascular stent adherence to the vascular wall after implantation.

Figure 10.

Figure 10

Open in a new tab

(a) Radical force of stents proximal in dependence of pressure and (b) radical force of stents.

3.4. Bionic Surface Structures and Morphology of the PTLG Materials by the SSD

After modification and enhancement of the PTLG random copolymer by the SSD process, the final product can be prepared by weaving, mechanical cutting, laser engraving, and other methods. For example, the SSD wire can be used to prepare monofilament sutures, the SSD wire can be woven into vascular scaffolds, the SSD tube can be mechanically carved into vascular scaffolds, the SSD cylindrical embryo can be mechanically cut into bone screws, and the SSD plate can be mechanically cut into bone fixation plates. BionX, a Finnish company, has applied PTLG random copolymer bone fixation plates prepared by SSD with highly oriented fibers where were then molded and sintered into shape in clinical practice.29−31 In addition, there have been reports of PTLG random copolymer bone screws prepared by SSD of PTLG random copolymer cylindrical embryos and mechanical cutting, which have better mechanical strength and degradation performance than the PTLG random copolymer bone screws formed by BionX through SSD fiber compression sintering.

As shown in Figure 11a, the prepared self-reinforcement PTLG thread products can be used as high-performance biodegradable surgical sutures. The tensile strength of the Mersilk suture used in clinical practice reaches 320 MPa,32 while the tensile strength of unmodified PLA was about 60 MPa. According to Figure 9, the tensile strength of the PTLG sample can reach 333 MPa. Therefore, the research method used in this paper to prepare PTLG thread products are expected to be used as monofilament sutures.

Figure 11.

Figure 11

Open in a new tab

(a) Self-reinforcing monofilament; (b) self-reinforcing vascular stent; (c) SEM images of PTLG before drawing; and (d) SEM images of PTLG-1200% samples.

One of the most important properties of vascular stents was their mechanical properties. Excellent mechanical properties enabled vascular stents to provide sufficient radial support for vascular stenosis, alleviate obstruction, and facilitate blood flow. If the vascular wall in the lesion area was thick or the lesion was severe, the vascular stent needed to provide greater radial support to expand the blood vessels and maintain smooth blood flow. The materials currently used in clinical practice for vascular stents such as stainless steel presented a strength of over 200 MPa. Ideal vascular stent materials have a strength of over 200 MPa, while PLA presented a strength of about 60 MPa.32Figure 9 indicates that the materials were prepared through the synergistic effect of PTLG and SSD. The mechanical strength and modulus of the material reached 333 MPa and 1.93 GPa, respectively, which met the requirements for the use of vascular stents from the perspective of material mechanical properties.

The SEM surface morphology observation results of PTLG-100% and PTLG-1200% samples are shown in Figure 11c,d, respectively. The initial sample PTLG-100% exhibited a smooth and flat surface structure. The surfaces with a bionic submicrometer morphology, the vessel-like bionic surface with parallel grooves and ridges ranging from 2 to 15 μm, were fabricated for PTLG-1200% samples by the SSD process. According to Figure 2, the SEM images indicated that the inner surface of the aortic intima was a ridge and groove structure along the blood flow direction, and the width of this regular structure was in the range 1–10 μm. We evaluated the similarity between the surface of PTLG samples and blood vessels on two aspects: overall morphology and the width and height of the parallel ridges and grooves. Besides, SEM images of the inner surface of blood vessels in Figure 2 and the surface of oriented PTLG obtained by SSD are shown in Figure 11d. By comparing Figures 2 and 11, the similarity in the overall appearance of the surface can be seen. Furthermore, based on the SEM characterization shown in Figures 2 and 11d, the three-dimensional morphology of the surface of the material by the SSD process was quantified. The width of the surface was measured to be about 2–15 μm, which was very similar to the structures on the inner surface of blood vessels that are 1–10 μm wide. In summary, the overall morphology and the spacing and height of the surface ridge structure of the material prepared by SSD were very similar to the one of internal blood vessels at the micro scale. The special bionic surface structure caused an increase of the interface area between the materials and cell to provide more binding sites for osteoblasts, which could be advantageous to the cell adsorption and adhesion in grooves by the mechanically interlocking effect.

3.5. Evaluation of Biocompatibilities of the PTLG Materials by the SSD Process

The CCK-8 assay kit was currently a commonly used method for detecting the cytotoxicity of biomaterials. Its reagents include WST-8 [2-(2-methoxy-4-nitrophenyl) −3-(4-nitro)-5-(2,4-disulfobenzoic)-2H-tetrazole monosodium salt], which was reduced by dehydrogenases in the cell mitochondria under the action of electron coupling agents to a highly water-soluble methylbenzene product. The number of methylbenzene products was directly proportional to the number of live cells and can indirectly reflect the number of live cells. CCK-8 toxicity testing was performed on PTLG samples subjected to SSD. Figure 12 shows the absorbance (O.D.) values at 2d, 4d, and 6d for PTLG samples with different drawing ratios tested by CCK-8. As the draw ratio increased, the O.D. of the samples increased. The values gradually increased, indicating that SSD effectively improved the cell compatibility. Among them, the PTLG samples with a maximum stretching ratio of 1200% has the highest O.D. The value indicated that it has the best cell compatibility among PTLG materials. The experimental results indicated that the PTLG vascular stent peripheral processed by SSD presented good cell compatibility.

Figure 12.

Figure 12

Open in a new tab

Cell viability by the CCK-8 kit after incubation with the PTLG stent for 2, 4, and 6 days.

The live/dead cell viability tests of PTLG samples with different draw ratios are shown in Figure 13. Calcein-AM can only pass through the living cell membrane and was cleaved by esterase to form an impermeable polar molecule, which was retained in the cell and emits strong green fluorescence. PI cannot penetrate the cell membrane of living cells but can only penetrate the disordered area of dead cell membranes and reach the nucleus, embedding in the double helix of deoxyribonucleic acid (DNA) to produce red fluorescence. As shown in Figure 13, there were a large number of green live cells on the surface of all PTLG samples, and only a small number of red dead cells were observed, indicating that the prepared PTLG samples have good cell compatibility after the SSD. As the draw ratio increases, the number of live cells on the surface of the samples gradually increased, while the number of dead cells decreased, indicating that the microgroove structure on the surface of the PTLG samples effectively improved the adsorption and proliferation of cells on the material surface. The results were consistent with the CCK-8 results, indicating that the formation of parallel microgroove structures on the PTLG samples surface effectively improved the adhesion, growth, and proliferation of cells on the material surface.

Figure 13.

Figure 13

Open in a new tab

Live/dead cell viability tests of PTLG samples with different draw ratios (green fluorescence represents live cells and red fluorescence represents dead cells).

The representative immunofluorescence microscopy images of the cell nucleus (blue) and cell cytoskeleton (green) staining of L929 cultured on the PTLG samples for 48 h are shown in Figure 14. Phalloidin can selectively bind to intracellular actin microfilaments (F-actin), and under an optical microscope, the fluorescence-labeled Phalloidin ye conjugate can clearly display the morphology and distribution of intracellular microfilaments. Hematoxylin dye can stain the cell nucleus. The phosphate groups on the two strands of DNA are outward, negatively charged, and acidic and easily bind with the positively charged hematoxylin basic dyes through ionic bonds to stain. Hematoxylin turns blue in alkaline solutions, so the cell nucleus was dyed blue. As the draw ratio increased, the number of stained nuclei gradually increased. In addition, as the draw ratio increased, the morphology of the cytoskeleton gradually spread and the cell antennae extended. The trend of changes in the nucleus and cytoskeleton with the draw ratio and the evolution of the shish kebab structure of the formation trend of the PTLG surface microgroove structures were basically consistent, which once again proved that the SSD simultaneously improved the mechanical properties and biocompatibility of materials. On one hand, the shish kebab crystal structure significantly improved the mechanical strength of the material, allowing PTLG samples to meet the requirements for mechanical properties of biomedical materials. On the other hand, this shish kebab structure exhibited submicron-level parallelly and orderly arranged micro-groove structures on the material surface, which effectively promoted cell adsorption and improved cell compatibility on the material surface. When biomaterials are implanted into tissues, they quickly react with cells, which is also an extended reaction after tissue damage, and the beginning of the reaction is when cells sense the extracellular mechanical signals. Cells perceive mechanical signals through various means, including structural proteins, mechanosensitive channels, enzymes, and receptors. Previous studies on traditional mechanosensitive channels have found that mechanosensitive channels are mostly located in nonexcitable cells, which is not entirely consistent with the theory that life evolution and information transmission require mechanosensitive channels. Costa et al. first discovered and identified Piezo1 mechanically sensitive ion channels and then discovered Piezeo2 through sequence homology, filling this gap. Piezo1 and Piezo2 genes are important components encoding mammalian mechanosensitive ion channels and have significant physiological and pathological functions. Piezo1 is highly expressed in multiple tissues such as the lungs, fat, and intestines and plays an important role in the development of touch, blood vessels and lymphatic vessels, blood pressure regulation, and red blood cell volume regulation. Piezo1 can respond to mechanical stimuli and cause ions to enter cells, thereby inducing cell excitation and signal transduction, converting mechanical stimuli such as touch, traction, and static pressure into chemical and electrical signals. Therefore, PTLG biomimetic interfaces with complex topological structures, through the characteristics of implant cell interface biomimetic structures, carry cells or actioned points in a nonplanar uniform manner, generate external mechanical induction, activate Piezo1 receptors in cells, regulate the expression of proteins and related genes in cells, and promote cell differentiation and proliferation.

Figure 14.

Figure 14

Open in a new tab

Representative immunofluorescence microscopy images of DAPI (blue) and phalloidin (green) staining of L929 cultured on the PTLG samples for 48 h.

4. Conclusions

The green and biodegradable materials with great mechanical properties and biocompatibility will offer new opportunities for next-generation, high-performance biological materials. Herein, the novelty oriented shish kebab crystals of a novel PTLG vascular stent are first reported to be successfully fabricated through a feasible SSD process to simultaneously enhance the mechanical performance and biocompatibility. The crystal structure of this self-reinforced vascular stent was transformed from spherulites to shish kebab crystals, which indicates a mechanical interlocking effect and prevents the lamellae from slipping with a significant improvement of mechanical strength to 333 MPa. Meanwhile, it is different from typical biomedical polymers with smooth surface structures, where the as-obtained PTLG vascular stent exhibits a bionic surface morphology with a parallel micro groove and ridge structure. These ridges and grooves were attributed to the reorganization of cytoskeleton fiber bundles following the direction of the blood flow shear stress. The structure and parameters of these morphologies were highly similar to the inner surface of blood vessels of humans, which facilitates cell adhesion growth to improve its proliferation, differentiation, and activity on the surface of PTLG.

Acknowledgments

This work was sponsored by the National Natural Science Foundation of China (nos. 52304141, 52374125, and 52104091), the Coal Mining and Designing Department, Tiandi Science & Technology Co Ltd (nos. 2023-TD-MS0013, 2019-TD-ZD006, and KJ-2022-KCZD-01), and the “Science and Technology Innovation Action Program” of Shanghai Science and Technology Commission (nos. 21ZR1412100 and no. 23ZR1480300).

Supporting Information Available

The Supporting Information is available free of charge at https://pubs.acs.org/doi/10.1021/acsami.3c18093.

  • Characterization of PTLG samples, mechanical properties of PTLG samples, and crystal morphology of PTLG samples (PDF)

The authors declare no competing financial interest.

Supplementary Material

am3c18093_si_001.pdf (280KB, pdf)

References

  1. Raquez J. M.; Habibi Y.; Murariu M.; Dubois P. Polylactide (PLA)-based nanocomposites. Prog. Polym. Sci. 2013, 38, 1504–1542. 10.1016/j.progpolymsci.2013.05.014. [DOI] [Google Scholar]
  2. Lim L. T.; Auras R.; Rubino M. Processing technologies for poly (lactic acid). Prog. Polym. Sci. 2008, 33, 820–852. 10.1016/j.progpolymsci.2008.05.004. [DOI] [Google Scholar]
  3. Rasal R. M.; Janorkar A. V.; Hirt D. E. Poly (lactic acid) modifications. Prog. Polym. Sci. 2010, 35, 338–356. 10.1016/j.progpolymsci.2009.12.003. [DOI] [Google Scholar]
  4. Tsuji H. Poly (lactic acid) stereocomplexes: A decade of progress. Adv. Drug Deliver. Rev. 2016, 107, 97–135. 10.1016/j.addr.2016.04.017. [DOI] [PubMed] [Google Scholar]
  5. Lasprilla A. J.; Martinez G. A.; Lunelli B. H.; Jardini A. L.; Filho R. M. Poly-lactic acid synthesis for application in biomedical devices-A review. Biotechnol. Adv. 2012, 30, 321–328. 10.1016/j.biotechadv.2011.06.019. [DOI] [PubMed] [Google Scholar]
  6. Jung Y.; Kim S. S.; Kim Y. H.; Kim S. H.; Kim B. S.; Kim S.; Choi C.; Kim S. A poly (lactic acid)/calcium metaphosphate composite for bone tissue engineering. Biomaterials 2005, 26, 6314–6322. 10.1016/j.biomaterials.2005.04.007. [DOI] [PubMed] [Google Scholar]
  7. Kao C. T.; Lin C. C.; Chen Y. W.; Ye h C. H.; Fang H. Y.; Shie M. Y. Poly (dopamine) coating of 3D printed poly (lactic acid) scaffolds for bone tissue engineering. Mater. Sci. Eng. C-Mater. 2015, 56, 165–173. 10.1016/j.msec.2015.06.028. [DOI] [PubMed] [Google Scholar]
  8. Majola A.; Vainionpää S.; Vihtonen K.; Mero M.; Vasenius J.; Törmälä P.; Rokkanen P. Absorption, biocompatibility, and fixation properties of polylactic acid in bone tissue: an experimental study in rats. Clin. Orthop. Relat. Res. 1991, 268, 260–269. [PubMed] [Google Scholar]
  9. Hou L. D.; Li Z.; Pan Y.; Sabir M.; Zheng Y. F.; Li L. A review on biodegradable materials for cardiovascular stent application. Front. Mater. Sci. 2016, 10, 238–259. 10.1007/s11706-016-0344-x. [DOI] [Google Scholar]
  10. Agrawal C. M.; Haas K. F.; Leopold D. A.; Clark H. G. Evaluation of poly (L-lactic acid) as a material for intravascular polymeric stents. Biomaterials 1992, 13, 176–182. 10.1016/0142-9612(92)90068-Y. [DOI] [PubMed] [Google Scholar]
  11. Schellhammer F.; Berlis A.; Bloss H.; Pagenstecher A.; Schumacher M. Poly-lactic-acid coating for endovascular stents: Preliminary results in canine experimental arteriovenous fistulae. Invest. Radiol. 1997, 32, 180–186. 10.1097/00004424-199703000-00008. [DOI] [PubMed] [Google Scholar]
  12. Wu Z.; Zhao J.; Wu W.; Wang P.; Wang B.; Li G.; Zhang S. Radial compressive property and the proof-of-concept study for realizing self-expansion of 3D printing polylactic acid vascular stents with negative Poisson’s ratio structure. Materials 2018, 11, 1357–1365. 10.3390/ma11081357. [DOI] [PMC free article] [PubMed] [Google Scholar]
  13. Dev A.; Binulal N. S.; Anitha A.; Nair S. V.; Furuike T.; Tamura H.; Jayakumar R. Preparation of poly (lactic acid)/chitosan nanoparticles for anti-HIV drug delivery applications. Carbohyd. Polym. 2010, 80, 833–838. 10.1016/j.carbpol.2009.12.040. [DOI] [Google Scholar]
  14. Yu L.; Liu H.; Xie F.; Chen L.; Li X. Effect of Annealing and Orientation on Microstructures and Mechanical Properties of Polylactic Acid. Polym. Eng. Sci. 2008, 48, 634–641. 10.1002/pen.20970. [DOI] [Google Scholar]
  15. Gupta B.; Revagade N.; Anjum N.; Atthoff B.; Hilborn J. Preparation of poly(lactic acid) fiber by dry-jet-wet-spinning. I. Influence of draw ratio on fiber properties. J. Appl. Polym. Sci. 2006, 100, 1239–1246. 10.1002/app.23497. [DOI] [Google Scholar]
  16. Lopes M. S.; Jardini A. L.; Filho R. M. Poly (lactic acid) production for tissue engineering applications. Procedia Eng. 2012, 42, 1402–1413. 10.1016/j.proeng.2012.07.534. [DOI] [Google Scholar]
  17. Santoro M.; Shah S. R.; Walker J. L.; Mikos A. G. Poly (lactic acid) nanofibrous scaffolds for tissue engineering. Adv. Drug. Deliver. Rev. 2016, 107, 206–212. 10.1016/j.addr.2016.04.019. [DOI] [PMC free article] [PubMed] [Google Scholar]
  18. Tripathi N.; Misra M.; Mohanty A. K. Durable polylactic acid (PLA)-based sustainable engineered blends and biocomposites: Recent developments, challenges, and opportunities. ACS Eng. Au 2021, 1, 7–38. 10.1021/acsengineeringau.1c00011. [DOI] [Google Scholar]
  19. Nofar M.; Sacligil D.; Carreau P. J.; Kamal M. R.; Heuzey M. C. Poly (lactic acid) blends: Processing, properties and applications. Int. J. Biol. Macromol. 2019, 125, 307–360. 10.1016/j.ijbiomac.2018.12.002. [DOI] [PubMed] [Google Scholar]
  20. Blasi P. Poly (lactic acid)/poly (lactic-co-glycolic acid)-based microparticles: An overview. J. Pharm. Investig. 2019, 49, 337–346. 10.1007/s40005-019-00453-z. [DOI] [Google Scholar]
  21. Fan T.; Ye W.; Du B.; Zhang Q.; Gong L.; Li J.; Lin S.; Fan Z.; Liu Q. Effect of Segment Structures on the Hydrolytic Degradation Behaviors of Totally Degradable Poly(L-lactic acid)-based Copolymers. J. Appl. Polym. Sci. 2019, 136, 47887. 10.1002/app.47887. [DOI] [Google Scholar]
  22. Fan T.; Qin J.; Meng X.; Li J.; Liu Q.; Wang G. Biodegradable Membrane of Poly (l-lactide acid-dioxanone-glycolide) and Stereocomplex Poly (lactide) with Enhanced Crystallization and Biocompatibility. Front. Bioeng. Biotechnol. 2022, 10, 1021218. 10.3389/fbioe.2022.1021218. [DOI] [PMC free article] [PubMed] [Google Scholar]
  23. Samantaray P. K.; Little A.; Haddleton D. M.; McNally T.; Tan B.; Sun Z.; Huang W.; Ji Y.; Wan C. Poly (glycolic acid)(PGA): A versatile building block expanding high performance and sustainable Bioplastic applications. Green Chem. 2020, 22, 4055–4081. 10.1039/d0gc01394c. [DOI] [Google Scholar]
  24. Feng P.; Shen S.; Shuai Y.; Peng S.; Shuai C.; Chen S. PLLA grafting draws GO from PGA phase to the interface in PLLA/PGA bone scaffold owing enhanced interfacial interaction. Sustain. Mater. Techno. 2023, 35, e00566 10.1016/j.susmat.2023.e00566. [DOI] [Google Scholar]
  25. Li J.; Chen Q.; Zhang Q.; Fan T.; Gong L.; Ye W.; Fan Z.; Cao L. Improving Mechanical Properties and Biocompatibilities by Highly Oriented Long Chain Branching Poly(lactic acid) with Bionic Surface Structures. ACS Appl. Mater. Interfaces 2020, 12, 14365–14375. 10.1021/acsami.9b20264. [DOI] [PubMed] [Google Scholar]
  26. Li J.; Ye W.; Fan Z.; Cao Lu. A Novel Stereocomplex Poly(lactic acid) with Shish-Kebab Crystals and Bionic Surface Structures as Bioimplant Materials for Tissue Engineering Applications. ACS Appl. Mater. Interfaces 2021, 13, 5469–5477. 10.1021/acsami.0c17465. [DOI] [PubMed] [Google Scholar]
  27. Liao T.; Zhao X.; Yang X.; Coates P.; Whiteside B.; Jiang Z.; Men Y. Structural Evolution of Flow-Oriented High Density Polyethylene Upon Heating: In Situ SAXS and WAXD Studies. Polymer 2019, 180, 121698. 10.1016/j.polymer.2019.121698. [DOI] [Google Scholar]
  28. Arima Y.; Iwata H. Effect of Wettability and Surface Functional Groups on Protein Adsorption and Cell Adhesion Using Well-Defined Mixed Self-Assembled Monolayers. Biomaterials 2007, 28, 3074–3082. 10.1016/j.biomaterials.2007.03.013. [DOI] [PubMed] [Google Scholar]
  29. Nakamura M.; Nagai A.; Hentunen T.; Salonen J.; Sekijima Y.; Okura T.; Hashimoto K.; Toda Y.; Monma H.; Yamashita K. Surface Electric Fields Increase Osteoblast Adhesion Through Improved Wettability on Hydroxyapatite Electret. ACS Appl. Mater. Interfaces 2009, 1, 2181–2189. 10.1021/am900341v. [DOI] [PubMed] [Google Scholar]
  30. Gehlen D. B.; De Lencastre Novaes L. C.; Long W.; Ruff A. J.; Jakob F.; Haraszti T.; Chandorkar Y.; Yang L.; van Rijn P.; Schwaneberg U.; De Laporte L.; De L. L. Rapid and Robust Coating Method to Render Polydimethylsiloxane Surfaces Cell-Adhesive. ACS Appl. Mater. Interfaces 2019, 11, 41091–41099. 10.1021/acsami.9b16025. [DOI] [PubMed] [Google Scholar]
  31. Tuomi J. T.; Björkstrand R. V.; Pernu M. L.; Salmi M. V. J.; Huotilainen E. I.; Wolff J. E. H.; Vallittu P. K.; Mäkitie A. A. In Vitro Cytotoxicity and Surface Topography Evaluation of Additive Manufacturing Titanium Implant Materials. J. Mater. Sci. 2017, 28, 53. 10.1007/s10856-017-5863-1. [DOI] [PubMed] [Google Scholar]
  32. McMahon S.; Bertollo N.; Cearbhaill E. D. O.; Salber J.; Pierucci L.; Duffy P.; Dürig T.; Bi V.; Wang W. Bio-resorbable polymer stents: a review of material progress and prospects. Prog. Polym. Sci. 2018, 83, 79–96. 10.1016/j.progpolymsci.2018.05.002. [DOI] [Google Scholar]

Associated Data

This section collects any data citations, data availability statements, or supplementary materials included in this article.

Supplementary Materials

am3c18093_si_001.pdf (280KB, pdf)

Articles from ACS Applied Materials & Interfaces are provided here courtesy of American Chemical Society

RESOURCES