- Email: [email protected]
Poly[2-(methacryloyloxy)ethyl choline phosphate] functionalized polylactic acid film with improved degradation resistance both in vitro and in vivo
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Contents lists available at ScienceDirect
Colloids and Surfaces B: Biointerfaces journal homepage: www.elsevier.com/locate/colsurfb
Poly[2-(methacryloyloxy)ethyl choline phosphate] functionalized polylactic acid film with improved degradation resistance both in vitro and in vivo Qiangwei Xina, Tong Suna, Min Lua, Zuxin Wangb, Keke Songb, Dong Hud,e, Zhiqiang Lid, Jun Luoc, Xingyu Chena,b,*, Jianshu Lic, Jie Wenga,b,* a Key Laboratory of Advanced Technologies of Materials (Ministry of Education), School of Materials Science and Engineering, Southwest Jiaotong University, Chengdu, 610031, China b College of Medicine, Southwest Jiaotong University, Chengdu, 610031, China c College of Polymer Science and Engineering, Sichuan University, Chengdu, 610065, China d Department of Orthopedics, The General Hospital of Western Theater Comand, Chengdu, 610083, China e School of Clinical Medicine, Chengdu Medicine College, Chengdu, 610500, China
A R T I C LE I N FO
A B S T R A C T
Keywords: Surface modification Degradation Surface-initiated atom transfer radical polymerization (SI-ATRP) Polylactic acid (PLA)
In this study, the surface-initiated atom transfer radical polymerization (SI-ATRP) of 2-(methacryloyloxy) ethyl choline phosphate (MCP) was successfully carried out via the ATRP initiator immobilized on the surfaces of polylactic acid (PLA) films. Different amounts of PMCP polymer brushes were constructed on the PLA surface to investigate the effects of the biological and degradation properties before and after modification. The results showed that the hydrophilicity of the surface of PLA were improved by MCP modification. In addition, there are no significant influence on the structure and crystallinity of the film before and after modification, except for the increased slightly thermal stability. Since the PMCP polymer brush forms a "protection" effect on the surface, the films showed an excellent property of resistant to hydrolysis even with obviously improved hydrophilicity. Furthermore, with the increase of the amount of introduced MCP monomer, the hydrophilicity and degradation resistance have been further improved. The in vivo animal experiment also verified this degradation resistance. Thereby, this strategy can be used to modulate the degradation rate of degradable polymers via surface modification.
1. Introduction Biodegradable polymers can effectively avoid the damage caused by secondary surgery after tissue healing. In addition, the rate of polymer degradation should be matched with the growth rate of the tissue to better meet the demands of tissue engineering [1,2]. Therefore, it is very necessary to prepare a polymer with a tunable degradation behavior. Polylactide (PLA) is an FDA approved biomedical polyester with excellent biocompatibility and biodegradability which can be metabolized in vivo [3]. It has been widely used for drug sustained release, surgical suture and orthopedic fixation and tissue repair materials owing to its excellent properties [4]. However, the biomedical applications of polylactide are somewhat restricted by its hydrophobicity and limited control of degradation [5–9]. Over the past decades, PLA has been modified to overcome
hydrophobic defects by various methods, including modifying the chain end-groups, copolymerization with hydrophilic monomers, control of the macromolecular architecture and surface modification, to improve its performance and make it better match the needs for tissue engineering after modification [10–13]. In addition, surface modification also can be a tool to adjust the degradation rate of degradable polymers [14]. PLA have been shown to degrade mainly by hydrolysis of the ester bond into lactic acid and can be removed from the body by normal metabolic pathway, the main factors affecting the degradation rate are: pH, temperature, and structure of the polymer [15–17]. Hydrophilic end groups such as the hydroxyl group can catalyze the degradation while the hydrophobic end groups such as methyl group can inhibit the catalysis. Therefore, people often introduce hydrophilic/hydrophobic end groups to adjust the degradation rate of the polymer. In general, the degradation rate of the material will increase as the hydrophilicity had been improved.
⁎ Corresponding authors at: Key Laboratory of Advanced Technologies of Materials (Ministry of Education), School of Materials Science and Engineering, Southwest Jiaotong University, Chengdu, 610031, China. E-mail addresses: [email protected] (X. Chen), [email protected] (J. Weng).
https://doi.org/10.1016/j.colsurfb.2019.110630 Received 2 October 2019; Received in revised form 4 November 2019; Accepted 6 November 2019 0927-7765/ © 2019 Elsevier B.V. All rights reserved.
Please cite this article as: Qiangwei Xin, et al., Colloids and Surfaces B: Biointerfaces, https://doi.org/10.1016/j.colsurfb.2019.110630
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 1. Schematic illustrattion of the surface modification process of PLA films by MCP.
300 μm, and the films were washed with an ethanol and deionized water mixture (1:1, v/v) for three times. The PLA films were dried at 50 °C for 2 h and cut into 2 cm × 2 cm for use in subsequent surface reactions.
Brooks et al. reported an interesting new zwitterionic molecule, i.e., choline phosphate (CP), with a reverse order of phosphate choline (PC) in cell membrane [18]. Recently, we synthesized CP into MCP, which can be designed into various polymeric molecular structures via SIATRP on the surface to improve the hydrophilicity of materials while endowing different performanc such as antifouling, cell adhesion promotion, and osteogenic differentiation properties [19–25]. However, to the best of our knowledge, there are no reports of the introduction of CP groups onto PLA surfaces, let alone further investigation on its degradation changes. The SI-ATRP method can be carried out under mild conditions to initiate surface grafting on different substrates [26–30]. In this work, we first prepared PLA films with ATRP initiation sites and then synthesized MCP monomers. After that, different amounts of MCP monomers were polymerized onto the surface of PLA films via SI-ATRP to form PMCP polymer brush with different molecular weights (Fig. 1). Finally, the effect of introduction of different proportions of PMCP polymer brush on the surface and degradation of the PLA films were investigated both in vitro and in vivo.
2.3. Immobilization of ATRP initiator on PLA film surface The prepared PLA films were immersed in 1, 6-hexanediamine/2propanol solution (60 mg/mL) for 6 h at 40 °C to introduce amino group on the PLA films, which were denoted as PLA-NH2 films. The ninhydrin analysis method was used to quantitatively detect the amount of NH2 groups on the PLA-NH2 films [31]. The PLA-NH2 films were then immersed in 2-bromoisobutyryl bromide (BIBB) solution which was composed of 1 mL BIBB and triethylamine mixed in 40 mL of anhydrous n-hexane. The reaction was first carried out at 0 °C for 2 h and then at 25 °C for another 24 h to obtain the bromo-terminated PLA surface (PLA-Br). Finally, the PLA-Br films were washed with an ethanol and deionized water mixture (1:1, v/v) for three times. 2.4. Surface-initiated ATRP of MCP
2. Experimental To prepare PLA-PMCP films, the reaction was carried out using [MCP (2 or 4 mmol)]: [CuBr]: [CuBr2]: [PMDETA] at the molar feed ratio of 100:2:1:0.2 in the methanol and deionized water mixture (5:1, v/v) as a solvent. The detailed procedure was adding CuBr2, PMDETA and monomeric MCP (2 or 4 mmol) to the round bottom flask, then performing the operations of "vacuum-argon" three times in the flask encircled ice. The CuBr was added quickly in the flask after venting with argon for 20 min, the operations of “vacuum-argon” was repeated three times and sealed tightly. After the reaction at 25 °C for 24 h, the PLA-PMCP2 (added 2 mmol monomeric MCP) and PLA-PMCP4 (added 4 mmol monomeric MCP) films were washed repeatedly with 50 % aqueous ethanol solution and dried [32,33].
2.1. Materials Poly (L-lactide) (PLLA, Mw = 150,000) was purchased from Jinan Daigang Biomaterial Co. Ltd. (Jinan, China). 1, 6-hexanediamine, 2propanol, methylalcohol (99.8 %), acetonitrile, 2-(dimethyl-amino) ethyl methacrylate (DMAEMA, 99.8 %), triethylamine (99.8 %), tetrahydrofuran (99.8 %), hydroquinol (99.0 %), CuBr2 (99.0 %), CuBr (99.0 %), sodium hydroxide and hydrochloric were all purchased from Haihong Regent Corporation. 1, 1, 4, 7, 7Pentamethyldiethylenetriamine (PMDETA, 98 %) was purchased from J &K Chemical Ltd. (Shanghai, China). 2-Bromoisobutyryl bromide (BIBB, 98 %) was purchased from Sigma (USA). 2-Chloro-1,3,2-dioxaphospholane-2-oxide (95.0 %) was purchased from Tokyo Chemical Industry. 2-(Methacryloyloxy) ethyl Choline Phosphate (MCP) was synthesized following our previous report [21]. Ultrapure water with resistivity of 18.2 MΩ · cm was used throughout.
2.5. Characterization The chemical composition of the PLA and PLA-PMCP films were determined by X-ray photoelectron spectroscopy (XPS, AXIS Ultra DLD, Kratos, Japan) with a focused monochromatic Al Kα source (1486.6 eV) for excitation [34]. The FTIR measurements were performed using an FTIR spectrometer (ATR-FTIR, PerkinElmer, Waltham, MA, USA). The surface topography and roughness of films were observed with atomic force microscopy (AFM, MFP-3D-BIO, Oxford Instrument, UK). The static water contact angles were used to test the hydrophilicity of the surfaces of these PLA films, using the sessile drop method on a contact
2.2. Preparation of PLA films PLA particles (3.75 g) were dissolved in 75 mL of anhydrous dichloromethane and mixed uniformly and defoamed by ultrasonic. Then, the solution was poured in three glass molds and volatilized at room temperature for 24 h. The thickness of the films was approximately 2
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 2. (A) XPS spectra of pristine PLA, PMCP immobilized PLA-PMCP films. Part of the graph (binding energy ranging from 0 to 450 eV) was enlarged for details. (B) High resolution XPS N 1s spectra and curve fitting of PLA-PMCP2 and PLA-PMCP4 films and P 2p core-level spectra of the PLA-PMCP2 and PLA-PMCP4 films, respectively.
2.7. Degradation studies in vivo
angle goniometer (DSA100, Gumbo Ham bur, Germany) [35]. The molecular weights and molecular weight distributions of the films were measured by gel permeation chromatography (GPC, Waters1424). The X-ray diffraction ((XRD, Dmax 1400, 40 kV, 110 mA, Japan) was used to detect the effect of modification on the crystal structure of the films. Thermogravimetric Analysis (TGAQ50 TA Instruments, New Castle, DE, USA) was used to detect the thermal properties of the films. Thermal analysis was measured by DSC (TGAQ50 TA Instruments, New Castle, DE, USA), and the percentage crystallinity (χc) was calculated using equation:
The animal experiments were approved by the Animal Research Committee of the University [37]. The in vivo degradation test of the PLA and PLA-PMCP films were carried out on 24 male SD rats (8 weeks, 200−220 g). The all films were cut into small discs with a diameter of 1 cm and subcutaneously immersed in 75 % alcohol for 1 h, and then sterilized for 1 h by ultraviolet light. The operating table and the surgical instruments were sterilized and disinfected before use. Rats were fasted for more than 12 h before the subcutaneous implantation, then anesthetized through peritoneal injection with 10 % chloral hydrate solution at 3.5 mL/kg. After shaving and disinfection, the sterilized PLA and PLA-PMCP films were implanted into the back where two perpendicular incisions were created. The wounds were closed with silk sutures and then caged. At 2, 4 and 8 weeks, the specimens were harvested. The harvested samples were weighed after dried, finally the SEM was used to detect the morphology of the degradation in vivo.
χc= ΔHf/ΔHf0 x 100 Where χc is the degree of crystallinity, ΔHf is the heat of fusion of the sample, and ΔHf0 is the heat of fusion of 100 % crystalline polymer. The value used for ΔHf0 was 93 J/g [36].
2.6. Degradation studies in vitro 3. Results and discusssion Polymer degradation behavior in different environments were quantified by dry weight loss. PLA and PLA-PMCP films (W0) were hydrolyzed in 10 mL PBS (0.1 M, pH = 7.3 ± 0.1)、hydrochloric acid (pH = 1) and sodium hydroxide (pH = 13) solutions at 37℃ with 110 rpm in water bath, respectively. The solutions were refreshed every 24 h. At each time point, recording the mass loss of the films after rinsed with ultra-pure water and dried, followed by weighing (Wt). The mass loss was calculated according to the following equation:
3.1. Surface modification and characteristics As shown in Fig. 1, the aminolysis reaction was carried out with a mixed solution of 1, 6-hexanediamine/isopropanol, which can break the lipid bonds on the surface of PLA to form amide bonds and free hydroxyl and amino groups [31,32,38]. The films containing hydroxyl and amino groups on the surface further reacted with 2-bromoisobutyryl bromide (BIBB) to obtain a macromolecular initiator PLA-Br, then initiating an ATRP reaction to obtain a PMCP polymer brush. It is noted that the ingredient was hydroxyl terminated polylactic acid (OH-PLAOH) which can enhance the reactivity in this work.
Weight loss (%) = [(W0-Wt)/W0] ×100 % Where, W0 is the initial weight and Wt is the residual weight of the dry films at time t. After hydrolyzed in sodium hydroxide (pH = 13) solutions at 4 h and 20 h, the surface morphologies of the PLA and PLA-PMCP films were observed by SEM.
3.1.1. Composition and content of surface chemical elements In order to evaluate the chemical composition of the surface of the 3
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 3. High resolution XPS C 1s spectra and curve fitting of pristine PLA, PLA-NH2 and N 1s core-level spectra, PLA-Br and Br 3d core-level spectra, PLA-PMCP2, and PLA-PMCP4 films, respectively.
PLA films after modification at various stages, XPS was used to detecte and calculate the elements of C, O, N, P, Br. XPS wide-scan spectra detected the obvious characteristic peaks of N 1s (about 400 eV), Br 3d (68.8), and P 2p (133.6 eV) on the surface of PLA-PMCP2 and PLAPMCP4 films after ATRP reaction (Fig. 2A), indicating that the PMCP polymer brush has been successfully grafted. Fig. 2B showed the N 1s core-level spectra of PLA-PMCP2 and PLA-PMCP4, there were two distinct signals at about 399.3 and 402.1 eV, among them the latter belongs to the signal of N+ which corresponding to quaternary amine salt (N+R4) [20] of CP groups for PLA-PMCP. The P 2p signal at BE of about 133.6 eV were detected after ATRP reactions, both [N+]/[N] and [P]/ [C] were increased from 0.62 to 0.78, 5.1 × 10-3 to 8.5 × 10-3 with the amount of added monomer MCP increased, respectively. Fig. 3 showed the C 1s core-level spectra of the original PLA, PLANH2, PLA-Br, PLA-PMCP2 and PLA-PMCP4 surface successively. The C 1s core-level spectra can be curve-fitted into three peak components with binding energies at 284.6, 286.4, and 288.7 eV in general, attributable to the CeH, CeO and O]CeO, respectively. However, the peak of O]CeN at about 287.0 eV was appeared on the surface of PLANH2 and PLA-Br with the reaction proceeds that associated with the linkages between the PLA film surface and ATRP initiators. In addition, the N 1s and Br 3d core-level spectra signal at BE of about 400 and 69 eV were characteristics of covalently bonded ammonia and bromine, respectively, the [N]/[C] ratio on PLA-NH2 surface is 4.3 × 10-2 and the [Br]/[C] ratio on PLA-Br surface is 5.5 × 10-3 (determined by sensitivity factor corrected N 1s and C 1s core horizontal area ratio) are consistent with previous studies [32,33]. What’s more, the appearance of CeN peak at 285.0 eV means that the successful grafting of PMCP onto PLA surface. In this work, the grafting yield (GY) of PMCP is defined as GY = (Wa - Wb)/A, where Wa and Wb are the weights of the dry film before and after modification, respectively, and A is the film area [32]. The GY is 11.3 and 27.1 μg/cm2 of PLA-PMCP2 and PLA-PMCP4, respectively. It can also see that the P content increased from 0.32 % to 0.52 % with the increased amount of adding MCP monomer, but the Br content in the initiation sites decreased gradually as shown in Table 1, because the PMCP polymer brush becomes longer and longer while the initiation
Table 1 XPS atomic concentrations of the chemical components of pristine PLA, PLANH2, PLA-Br, PLA-PMCP2 and PLA-PMCP4 films. Samples
C
O
N
Br
P
PLA PLA-NH2 PLA-Br PLA-PMCP2 PLA-PMCP4
65.79 61.93 61.66 63.33 61.63
34.21 35.41 36.70 33.37 33.63
0 2.66 1.30 2.89 1.27
0 0 0.34 0.09 0.05
0 0 0 0.32 0.52
sites was entraped within the polymeric chains. Also, the oxygen percentage did not significantly change during the surface modification while the carbon was noticeably reduced compared to pristine PLA. In addition, the changes of functional group had been detected by ATRFTIR. As shown in Fig. S1, the peak intensity of the C]O stretching vibration at 1749 cm−1 is gradually increased with the increase of surface polymer brush. These results also indicated that the PMCP has been grafted successfully on the PLA surface by SI-ATRP.
3.1.2. Determination of amino groups The ninhydrin method was used to qualitatively and quantitatively detect the existence of NH2 groups on the aminolyzed PLA films. As presented in Fig. S2a, the absorbance after the ninhydrin reacted with 1,4-dioxane–isopropanol (1:1, v/v) solution containing 1,6-hexanediamine of known concentration in the range of 450−650 nm and the maximum absorption peak after the color reaction of the aminolyzed films with ninhydrin at 560 nm. The standard curve: y = 0.106x + 0.096, R2 = 0.9986 obtained by plotting the absorbance (y) versus amino concentration (x) in 560 nm and the absorbance of aminolyzed PLA films reacted with Ninhydrin at 560 nm was shown in Fig. S2b. Through these experiments, we determined that when the membrane area is fixed and the concentration of 1,6 hexanediamine is 3.34 × 10-2 mg/mL, the calculated amino density is 2.2 × 10-7 mol/cm2. Calculation method: ρ(NH2) = N x V/S, where N is the concentration of the amine groups on the aminolyzed PLA films, which was obtained by referring to the standard curve when the ninhydrin is colored and 4
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 4. Surface morphology of PLA and PLA- PMCP films.
dissolved in the mixed solvent, V is the volume of the color developing solution and S is the surface area of aminolyzed PLA films.
hydrophilicity of these surfaces demonstrated that the grafting reaction was successful.
3.1.3. Surface mopography and roughness Fig. 4 and Table S1 showed the statistical results of two-dimensional and three-dimensional topography and roughness of samples before and after modification. As can be seen, the surface of the initial PLA film was relatively smooth, and the surface was more roughness when the PMCP polymer was grafted. Meanwhile, as the amount of MCP monomer input increased, the molecular weight of modified films increased slightly (measured by GPC, shown in Fig. S3 and Table S2), and lead the surface roughness also increased due to the addition of PMCP brushes.
3.2. Effect of surface modification on its structure and properties
3.1.4. Water contact angles The hydrophilicity of the surface is also changed during the modification. As shown in Fig. 5, compared with the pristine PLA, the static water contact angle of PLA-NH2 decreased from about 87.0° to 68.5°. After introducing BIBB onto the PLA-NH2 surface, the water contact angle increased slightly because of the hydrophobicity of the initiator. With the surface ATRP reaction introduced onto the PLA-Br surface, the water contact angle sustainably decreased to 61.0° and 51.5° for PLAPMCP2 and PLA-PMCP4, respectively. The hydrophilicity of the surface of the PLA films were improved obviously after introduced the PMCP polymer brushes, and the more polymer brushes the higher hydrophilicity. In summary, the results of XPS and the change of
Some previous studies on the hydrolytic degradation of poly (lactic acid) suggest that PLA degrades faster in alkaline environments [39]. The catalysts act to accelerate the degradation of PLA, the alkali or lipase usually added to promote degradation. The proteinase K is the most widely used enzyme which can degrade PLA, but proteinase K can only degrade non-crystalline PLLA segments, therefore, the addition of alkali is generally used for accelerated the degradation of PLLA. As shown in Fig. 6A, the degradation studies in phosphate buffer solutions (100 mM, pH = 7.3 ± 0.1) and hydrochloric acid (pH = 1) showed slight weight loss (barely degraded) after incubating samples at 37℃ for three months, while the degradation in NaOH solutions (pH = 13) just need several days for complete degradation. The hydrolytic degradation rate of original PLA films in all conditions was higher than that of PLAPMCP2 and PLA-PMCP4 films, even though they are more hydrophilic. For instance, the pristine PLA films have degraded 73.6 % after 20 h, while PLA-PMCP2 and PLA-PMCP4 degraded 36.4 % and 23.5 % in NaOH solutions, respectively. Moreover, it has been reported that the PLA film has different degradation mechanisms in several solutions, such as the hydrolysis is mainly the surface erosion in alkaline solution and noumenon erosion in phosphate buffer and acidic solution. The durability in acidic solution is similar to neutral conditions and higher than in alkaline environments [40,41]. These results are also consistent with our previous work about PMCP modified polycaprolactone (PCL) as a tissue engineering material [21]. It is well known that for partially crystallized PLLA, water first penetrates into the amorphous region, causing the ester bond to break, and when most of the amorphous region has degraded, the edge begins to degrade toward the center of the crystalline region and it is driven via nucleophilic attack of hydroxide ions on carbonyl carbon [42,43]. Therefore, there are several possible difficulties when the modified films degraded in solutions: (1) the zwitterion first forms a hydration layer with water molecules, making
The XRD, DSC and TG had been used to detect the effect of modification on the crystal structure and thermal properties of the PLA and PLA-PMCP films. As the results showed in Fig. S4, there were no significant influence on the structure and crystallinity of the films before and after modification. 3.3. Degradation study in vitro
Fig. 5. Static water contact angles of Pristine PLA, PLA-NH2, PLA-Br, PLAPMCP2 and PLA-PMCP4 films. 5
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 6. (A) Degradation profiles of pristine PLA, PLA-PMCP2 and PLA-PMCP4 films at 37℃ in PBS (pH = 7.3), hydrochloric acid (pH = 1), NaOH solution (pH = 13) for 20 h and degraded completely and (B) SEM images of pristine PLA, PLA-PMCP2 and PLA-PMCP4 films, before and after degradation in NaOH solution for 4 h and 20 h, respectively.
6
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 7. (A) Biodegradation profiles of PLA and PLA-PMCP films in vivo and (B) SEM images of pristine PLA, PLA-PMCP2 and PLA-PMCP4 films, implanted in vivo for 2 weeks, 4 weeks and 8 weeks, respectively.
in C content and a decrease in O content.
the water molecules less permeable; (2) there are less ester bond exposed on the film surface as they are covered by zwitterionic CP groups; (3) the slightly increased molecular weight of the PLA films after PMCP modification. As a result, the PMCP polymer brush grafted on the surface has a certain “protective” effect on the surface which weakens the nucleophilic attack of the OH- ion on the ester bond, thereby delaying the degradation rate. Furthermore, the more the grafted PMCP, the degradation rate of PLA-PMCP films are more tardiness. Scanning electron microscopy (SEM) was used to observe the surface morphology of modified PLA films before and after degradation. As shown in Fig. 6B, the PLA, PLA-PMCP2 and PLA-PMCP4 films were smooth before degradation, however, as the degradation time prolonged, a large number of gaps and pores began to appear on the surface of the PLA films, while the modified PLA- PMCP2 and PLA-PMCP4 were relatively intact. The SEM results showed that the erosion rate of PLA-PMCP4 was lower than the PLA-PMCP2 and PLA under the same conditions. Among them, PLA-PMCP4 film exhibited the best degradation resistance in terms of morphology, which also corresponds to the degradation weight loss results shown in Fig. 6A. These results further confirmed that the introduction of hydrophilic MCP onto the surface of PLA can improve its hydrophilicity while delay the degradation rate at the same time. In this study, the mechanism of resist degradation also be discussed using the chemical derivatization technique in XPS and analysed on the PLA-PMCP4 films and degradation in NaOH for 4 h and 20 h, respectively. As the C 1s core-level spectra of the films showed in Fig. S5, the peak intensity of CeO and O]CeO were decreased gradually compared with pristine PLA-PMCP4 film, which means the ester group fracture and the degradation of the films. Fig. S6 showed the N 1s corelevel spectra of the modified films after degradation in over time. The results showed that with the prolongation of degradation time, the proportion of [N+]/[N] on the surface of the modified films were decreasing, which further confirms the surface protection of PMCP and delayed its degradation. It can also see in Table S3 that as the degradation progressed, there was no significant difference between the ratio of C and O on the surface compared with pristine PLA-PMCP4 films at 4 h, but at 20 h, the more decrease of [N+]/[N] aggravated the degradation tendency of the films, and the content of C was increased obviously from 63.39 %–70.40 %, while the content of O gradually decreased from 34.67 % to 29.13 %, respectively. In summary, due to the protective effect of the PMCP polymer brush on the PLA surface, the degradation of PLA had been delayed for a period of time, when the protection of PMCP fails, the degradation of PLA will lead to an increase
3.4. Degradation studies in vivo Rats subsided the anesthesia when the pristine PLA and PLA-gPMCP films were implanted for 2 h and there was no obvious swelling or bleeding. After 24 h, they began to eat and walk without any rat death. All of them were treated with neck removal during the planned time. A representative macroscopic image of the surgical wound of the rats at different times after implantation of the films are shown in Fig. S7. At 2 weeks, the incisions healed substantially, after 4 and 8 weeks the incisions were completely healed and hardly detectable. The hair of the rats gradually grew and covered the surgical area. In addition, there was no significant difference in the macroscopic morphology of all the films before and after implantation at various times. The in vivo biodegradation profiles of pristine PLA and PLA-PMCP films were evaluated by measuring weight loss (%) over time (Fig. 7A). There were both chemical and enzymatic degradation of PLA in vivo, and the degradation rate in vivo was faster than that in vitro, which was similar to other polyester materials [44–46]. The mass loss of the films was faster in the first two weeks, and then the change was relatively flat and slightly different from in vitro degradation, which may be due to the metabolism of the body and the complex physiological environment in vivo. However, the PLA-PMCP films also exhibited slower degradation in vivo than that of PLA film, which was similar to the degradation in vitro, and that could provide a relatively stable structure at the beginning of implantation, which would be very meaningful in tissue engineering. The SEM images of the films before and after implantation in vivo have showed in Fig. 7B. We observed that the surfaces of all films are relatively smooth and there was no significant difference before implantation. As time went on, the surface of the PLA films began to crack at 2 weeks, while the morphology of the PLA-PMCP films still integrity. The cracks on the surface of PLA film had continued to expand, and tiny cracks also appeared on the PLA-PMCP films with the increase of implantation time. When at 8 weeks, the surface of PLA consisted of many small pieces, PLA-PMCP films were also increased a lot of cracks. In addition, the surface of the PLA film had the most cracks, followed by PLA-PMCP2 and PLA-PMCP4 films which indicated that the PLA film had the most degradation and the PLA-PMCP film had the best degradation resistance, which was consistent with the trend of mass loss in vivo. In summary, the PLA-PMCP films had a slower degradation rate than PLA film in vitro and in vivo which have exhibited excellent 7
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
properties of degradation resistance.
[16] L. Xu, K. Crawford, C.B. Gorman, Effects of temperature and pH on the degradation of poly (lactic acid) brushes, Macromolecules 44 (2011) 4777–4782. [17] P. Gunatillake, R. Mayadunne, R. Adhikari, Recent developments in biodegradable synthetic polymers, Biotechnol. Annu. Rev. 12 (2006) 301–347. [18] X. Yu, Z. Liu, J. Janzen, I. Chafeeva, S. Horte, W. Chen, R.K. Kainthan, J.N. Kizhakkedathu, D.E. Brooks, Polyvalent choline phosphate as a universal biomembrane adhesive, Nat. Mater. 11 (2012) 468–476. [19] X. Chen, T. Chen, Z. Lin, X. Li, W. Wu, J. Li, Choline phosphate functionalized surface: protein-resistant but cell-adhesive zwitterionic surface potential for tissue engineering, Chem. Commun. (Camb.) 51 (2015) 487–490. [20] Y. Jiang, Y. Su, L. Zhao, F. Meng, Q. Wang, C. Ding, J. Luo, J. Li, Preparation and antifouling properties of 2-(meth-acryloyloxy) ethyl cholinephosphate based polymers modified surface with different molecular architectures by ATRP, Colloids Surf. B 156 (2017) 87–94. [21] X. Chen, Z. Lin, Y. Feng, H. Tan, X. Xu, J. Luo, J. Li, Zwitterionic PMCP-modified polycaprolactone surface for tissue engineering: antifouling, cell adhesion promotion, and osteogenic differentiation properties, Small (2019) e1903784. [22] F. Huang, C. Ding, J. li, Resisting protein but promoting cell adhesion by choline phosphate: a comparative study with phosphorylcholine, J. Bioresour. Bioprod. 3 (2018) 3–8. [23] W. Wang, B. Wang, S. Liu, X. Shang, X. Yan, Z. Liu, X. Ma, X. Yu, Bioreducible polymer nanocarrier based on multivalent choline phosphate for enhanced cellular uptake and intracellular delivery of doxorubicin, ACS Appl. Mater. Interfaces 9 (2017) 15986–15994. [24] X. Chen, H. Shang, S. Cao, H. Tan, J. Li, A zwitterionic surface with general celladhesive and protein-resistant properties, RSC Adv. 5 (2015) 76216–76220. [25] X. Chen, M. Yang, B. Liu, Z. Li, H. Tan, J. Li, Multilayer choline phosphate molecule modified surface with enhanced cell adhesion but resistance to protein adsorption, Langmuir 33 (2017) 8295–8301. [26] Z. Wu, H. Chen, X. Liu, Y. Zhang, D. Li, H. Huang, Protein adsorption on poly(Nvinylpyrrolidone)-modified silicon surfaces prepared by surface-initiated atom transfer radical polymerization, Langmuir 5 (2009) 2900–2906. [27] G. Cheng, Z. Zhang, S. Chen, J.D. Bryers, S. Jiang, Inhibition of bacterial adhesion and biofilm formation on zwitterionic surfaces, Biomaterials 28 (2007) 4192–4199. [28] W.J. Yang, T. Cai, K.G. Neoh, E.T. Kang, G.H. Dickinson, S.L. Teo, D. Rittschof, Biomimetic anchors for antifouling and antibacterial polymer brushes on stainless steel, Langmuir 27 (2011) 7065–7076. [29] D.J. Siegwart, J.K. Oh, K. Matyjaszewski, ATRP in the design of functional materials for biomedical applications, Prog. Polym. Sci. 37 (2012) 18–37. [30] Q. Yang, J. Tian, M. Hu, Z. Xu, Construction of a comb-like glycosylated membrane surface by a combination of UV-induced graft polymerization and surface-initiated ATRP, Langmuir 23 (2007) 6684–6690. [31] Y. Zhu, C. Gao, X. Liu, T. He, J. Shen, Immobilization of Biomacromolecules onto aminolyzed poly (L-lactic acid) toward acceleration of endothelium regeneration, Tissue Eng. 10 (2004) 53–61. [32] F. Xu, Z. Wang, W. Yang, Surface functionalization of polycaprolactone films via surface-initiated atom transfer radical polymerization for covalently coupling celladhesive biomolecules, Biomaterials 31 (2010) 3139–3147. [33] F. Xu, X. Yang, C. Li, W. Yang, Functionalized polylactide film surfaces via surfaceinitiated ATRP, Macromolecules 44 (2011) 2371–2377. [34] M. Wan, H. Chen, Q. Wang, Q. Niu, P. Xu, Y. Yu, T. Zhu, C. Mao, J. Shen, Bioinspired nitric-oxide-driven nanomotor, Nat. Commun. 10 (2019) 966. [35] M. Wan, T. Xu, B. Chi, M. Wang, Y. Huang, Q. Wang, T. Li, W. Yan, H. Chen, P. Xu, C. Mao, B. Zhao, J. Shen, H. Xu, D. Shi, A safe and efficient strategy for the rapid elimination of blood lead in vivo based on a capture-fix-separate mechanism, Angew. Chem. Int. Ed. 58 (2019) 10582–10586. [36] M. Yasuniwa, K. Sakamo, Y. Ono, W. Kawahara, Melting behavior of poly (l-lactic acid): X-ray and DSC analyses of the melting process, Polymer 49 (2008) 1943–1951. [37] M. Wan, J. Zhang, Q. Wang, S. Zhan, X. Chen, C. Mao, Y. Liu, J. Shen, In situ growth of mesoporous silica with drugs on titanium surface and its biomedical applications, ACS Appl. Mater. Interfaces 9 (2017) 18609–18618. [38] Y. Gong, Y. Zhu, Y. Liu, Z. Ma, C. Gao, J. Shen, Layer-by-layer assembly of chondroitin sulfate and collagen on aminolyzed poly (L-lactic acid) porous scaffolds to enhance their chondrogenesis, Acta Biomater. 3 (2007) 677–685. [39] S.J.D. Jong, E.R. Ariasa, D.T.S. Rijkersb, C.F.V. Nostruma, J.J.K.V.D. Boschc, W.E. Hennink, New insights into the hydrolytic degradation of poly (lactic acid): participation of the alcohol terminus, Polymer 42 (2001) 2795–2802. [40] H. Tsuji, Y. Ikada, Properties and morphology of poly (L-lactide). 4. Effects of structural parameters on long-term hydrolysis of poly (L-lactide) in phosphatebuffered solution, Polym. Degrad. Stab. 67 (2000) 179–189. [41] H. Tsuji, Y. Ikada, Properties and morphology of poly (L-lactide). II. Hydrolysis in alkaline solution, Polym. Sci. Part A: Polym. Chem. 36 (1998) 59–66. [42] H. Fukuzaki, M. Yoshida, M. Asano, M. Kumakura, Synthesis of copoly (D, L-lactic acid) with relatively low molecular weight and in vitro degradation, Eur. Polym. J. 25 (1989) 1019–1026. [43] C.Y. Tham, Z.A.A. Hamid, Z.A. Ahmad, H. Ismail, Surface engineered poly (lactic acid) (PLA) microspheres by chemical treatment for drug delivery system, Key Eng. Mater. 594–595 (2013) 214–218. [44] J.M. Anderson, A. Rodriguez, D.T. Chang, Foreign body reaction to biomaterials, Semin. Immunol. 20 (2008) 86–100. [45] E.M. Christenson, J.M. Anderson, A. Hiltner, Oxidative mechanisms of poly (carbonate urethane) and poly (ether urethane) biodegradation: in vivo and in vitro correlations, J. Biomed. Mater. Res. 70 (2004) 245–255. [46] J.P. Santerre, K. Woodhouse, G. Laroche, R.S. Labow, Understanding the biodegradation of polyurethanes: from classical implants to tissue engineering materials, Biomaterials 26 (2005) 7457–7470.
4. Conclusions In the present study, different proportions of PMCP brush were successfully grafted on the PLA film surfaces via SI-ATRP. The modified films showed more resistance to degradation even though the hydrophilicity was improved obviously owing to the “protective” of PMCP polymer brush compared with pristine PLA. Moreover, with more grafted PMCP monomer, the PLA-PMCP films exhibited higher hydrophilicity and better degradation resistance. Meanwhile, the modified films in vivo also exhibited similar degradation resistance and a controlled rate of degradation in vitro environment, thereby suggesting it have a good application prospect in tissue engineering. Declaration of Competing Interest The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. Acknowledgements This work was supported by the National Natural Science Foundation of China (51803173, 51572228), Sichuan Province Science and Technology Innovation MiaoZi Funding (2018002), and Fundamental Research Funds for the Central Universities (2682019CX64). Appendix A. Supplementary data Supplementary material related to this article can be found, in the online version, at doi:https://doi.org/10.1016/j.colsurfb.2019.110630. References [1] L.S. Nair, C.T. Laurencin, Biodegradable polymers as biomaterials, Prog. Polym. Sci. 32 (2007) 762–798. [2] Y. Cheng, S. Deng, P. Chen, R. Ruan, Polylactic acid (PLA) synthesis and modifications: a review, Front. Chem. China 4 (2009) 259–264. [3] R. Kulkarni, K. Pani, C. Neuman, F. Leonard, Polylactic acid for surgical implants, Arch Surg. 93 (1966) 839–843. [4] B. Tyler, D. Gullotti, A. Mangraviti, T. Utsuki, H. Brem, Polylactic acid (PLA) controlled delivery carriers for biomedical applications, Adv. Drug Deliv. Rev. 107 (2016) 163–175. [5] N. Rotter, F. Ung, A.K. Roy, M. Vacanti, R.D. Eavey, C.A. Vacanti, L.J. Bonassar, Role for interleukin 1α in the inhibition of chondrogenesis in autologous implants using polyglycolic acid–polylactic acid scaffolds, Tissue Eng. Part A 11 (2005) 192–200. [6] T. Kissel, Y. Li, F. Unger, ABA-triblock copolymers from biodegradable polyester Ablocks and hydrophilic poly (ethylene oxide) B-blocks as a candidate for in situ forming hydrogel delivery systems for proteins, Adv. Drug Deliv. Rev. 54 (2002) 99–134. [7] M. Vert, Aliphatic polyesters: great degradable polymers that cannot do everything, Biomacromolecules 6 (2005) 538–546. [8] A.C. Albertsson, I.K. Varma, Recent developments in ring opening polymerization of lactones for biomedical applications, Biomacromolecules 4 (2003) 1466–1486. [9] E.S. Place, J.H. George, C.K. Williams, M.M. Stevens, Synthetic polymer scaffolds for tissue engineering, Chem. Soc. Rev. 38 (2009) 1139–1151. [10] M. Tang, A.F. Haider, C. Minelli, M.M. Stevens, C.K. Williams, Lactide polymerization co-initiated by carbohydrate esters and pyranoses, J. Polym. Sci. Part A: Polym. Chem. 46 (2008) 4352–4362. [11] X. Jiang, M.R. Smith, G.L. Baker, Water-soluble thermoresponsive polylactides, Macromolecules 41 (2008) 318–324. [12] C. Gottschalk, F. Wolf, H. Frey, Multi-arm star poly(L-lactide) with hyperbranched polyglycerol core, Macromol. Chem. Phys. 208 (2007) 1657–1665. [13] J. Ji, H. Zhu, J. Shen, Surface tailoring of poly (dl-lactic acid) by ligand-tethered amphiphilic polymer for promoting chondrocyte attachment andgrowth, Biomaterials 25 (2004) 1859–1867. [14] A. Hoglund, M. Hakkarainen, U. Edlund, A.C. Albertsson, Surface modification changes the degradation process and degradation product pattern of polylactide, Langmuir 26 (2010) 378–383. [15] A.M. Reed, D.K. Gilding, Biodegradable polymers for use in surgery poly (glycolic) / poly (lactic acid) homo and copolymers: 2. In vitro degradation, Polymer 22 (1980) 494–499.
8
Contents lists available at ScienceDirect
Colloids and Surfaces B: Biointerfaces journal homepage: www.elsevier.com/locate/colsurfb
Poly[2-(methacryloyloxy)ethyl choline phosphate] functionalized polylactic acid film with improved degradation resistance both in vitro and in vivo Qiangwei Xina, Tong Suna, Min Lua, Zuxin Wangb, Keke Songb, Dong Hud,e, Zhiqiang Lid, Jun Luoc, Xingyu Chena,b,*, Jianshu Lic, Jie Wenga,b,* a Key Laboratory of Advanced Technologies of Materials (Ministry of Education), School of Materials Science and Engineering, Southwest Jiaotong University, Chengdu, 610031, China b College of Medicine, Southwest Jiaotong University, Chengdu, 610031, China c College of Polymer Science and Engineering, Sichuan University, Chengdu, 610065, China d Department of Orthopedics, The General Hospital of Western Theater Comand, Chengdu, 610083, China e School of Clinical Medicine, Chengdu Medicine College, Chengdu, 610500, China
A R T I C LE I N FO
A B S T R A C T
Keywords: Surface modification Degradation Surface-initiated atom transfer radical polymerization (SI-ATRP) Polylactic acid (PLA)
In this study, the surface-initiated atom transfer radical polymerization (SI-ATRP) of 2-(methacryloyloxy) ethyl choline phosphate (MCP) was successfully carried out via the ATRP initiator immobilized on the surfaces of polylactic acid (PLA) films. Different amounts of PMCP polymer brushes were constructed on the PLA surface to investigate the effects of the biological and degradation properties before and after modification. The results showed that the hydrophilicity of the surface of PLA were improved by MCP modification. In addition, there are no significant influence on the structure and crystallinity of the film before and after modification, except for the increased slightly thermal stability. Since the PMCP polymer brush forms a "protection" effect on the surface, the films showed an excellent property of resistant to hydrolysis even with obviously improved hydrophilicity. Furthermore, with the increase of the amount of introduced MCP monomer, the hydrophilicity and degradation resistance have been further improved. The in vivo animal experiment also verified this degradation resistance. Thereby, this strategy can be used to modulate the degradation rate of degradable polymers via surface modification.
1. Introduction Biodegradable polymers can effectively avoid the damage caused by secondary surgery after tissue healing. In addition, the rate of polymer degradation should be matched with the growth rate of the tissue to better meet the demands of tissue engineering [1,2]. Therefore, it is very necessary to prepare a polymer with a tunable degradation behavior. Polylactide (PLA) is an FDA approved biomedical polyester with excellent biocompatibility and biodegradability which can be metabolized in vivo [3]. It has been widely used for drug sustained release, surgical suture and orthopedic fixation and tissue repair materials owing to its excellent properties [4]. However, the biomedical applications of polylactide are somewhat restricted by its hydrophobicity and limited control of degradation [5–9]. Over the past decades, PLA has been modified to overcome
hydrophobic defects by various methods, including modifying the chain end-groups, copolymerization with hydrophilic monomers, control of the macromolecular architecture and surface modification, to improve its performance and make it better match the needs for tissue engineering after modification [10–13]. In addition, surface modification also can be a tool to adjust the degradation rate of degradable polymers [14]. PLA have been shown to degrade mainly by hydrolysis of the ester bond into lactic acid and can be removed from the body by normal metabolic pathway, the main factors affecting the degradation rate are: pH, temperature, and structure of the polymer [15–17]. Hydrophilic end groups such as the hydroxyl group can catalyze the degradation while the hydrophobic end groups such as methyl group can inhibit the catalysis. Therefore, people often introduce hydrophilic/hydrophobic end groups to adjust the degradation rate of the polymer. In general, the degradation rate of the material will increase as the hydrophilicity had been improved.
⁎ Corresponding authors at: Key Laboratory of Advanced Technologies of Materials (Ministry of Education), School of Materials Science and Engineering, Southwest Jiaotong University, Chengdu, 610031, China. E-mail addresses: [email protected] (X. Chen), [email protected] (J. Weng).
https://doi.org/10.1016/j.colsurfb.2019.110630 Received 2 October 2019; Received in revised form 4 November 2019; Accepted 6 November 2019 0927-7765/ © 2019 Elsevier B.V. All rights reserved.
Please cite this article as: Qiangwei Xin, et al., Colloids and Surfaces B: Biointerfaces, https://doi.org/10.1016/j.colsurfb.2019.110630
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 1. Schematic illustrattion of the surface modification process of PLA films by MCP.
300 μm, and the films were washed with an ethanol and deionized water mixture (1:1, v/v) for three times. The PLA films were dried at 50 °C for 2 h and cut into 2 cm × 2 cm for use in subsequent surface reactions.
Brooks et al. reported an interesting new zwitterionic molecule, i.e., choline phosphate (CP), with a reverse order of phosphate choline (PC) in cell membrane [18]. Recently, we synthesized CP into MCP, which can be designed into various polymeric molecular structures via SIATRP on the surface to improve the hydrophilicity of materials while endowing different performanc such as antifouling, cell adhesion promotion, and osteogenic differentiation properties [19–25]. However, to the best of our knowledge, there are no reports of the introduction of CP groups onto PLA surfaces, let alone further investigation on its degradation changes. The SI-ATRP method can be carried out under mild conditions to initiate surface grafting on different substrates [26–30]. In this work, we first prepared PLA films with ATRP initiation sites and then synthesized MCP monomers. After that, different amounts of MCP monomers were polymerized onto the surface of PLA films via SI-ATRP to form PMCP polymer brush with different molecular weights (Fig. 1). Finally, the effect of introduction of different proportions of PMCP polymer brush on the surface and degradation of the PLA films were investigated both in vitro and in vivo.
2.3. Immobilization of ATRP initiator on PLA film surface The prepared PLA films were immersed in 1, 6-hexanediamine/2propanol solution (60 mg/mL) for 6 h at 40 °C to introduce amino group on the PLA films, which were denoted as PLA-NH2 films. The ninhydrin analysis method was used to quantitatively detect the amount of NH2 groups on the PLA-NH2 films [31]. The PLA-NH2 films were then immersed in 2-bromoisobutyryl bromide (BIBB) solution which was composed of 1 mL BIBB and triethylamine mixed in 40 mL of anhydrous n-hexane. The reaction was first carried out at 0 °C for 2 h and then at 25 °C for another 24 h to obtain the bromo-terminated PLA surface (PLA-Br). Finally, the PLA-Br films were washed with an ethanol and deionized water mixture (1:1, v/v) for three times. 2.4. Surface-initiated ATRP of MCP
2. Experimental To prepare PLA-PMCP films, the reaction was carried out using [MCP (2 or 4 mmol)]: [CuBr]: [CuBr2]: [PMDETA] at the molar feed ratio of 100:2:1:0.2 in the methanol and deionized water mixture (5:1, v/v) as a solvent. The detailed procedure was adding CuBr2, PMDETA and monomeric MCP (2 or 4 mmol) to the round bottom flask, then performing the operations of "vacuum-argon" three times in the flask encircled ice. The CuBr was added quickly in the flask after venting with argon for 20 min, the operations of “vacuum-argon” was repeated three times and sealed tightly. After the reaction at 25 °C for 24 h, the PLA-PMCP2 (added 2 mmol monomeric MCP) and PLA-PMCP4 (added 4 mmol monomeric MCP) films were washed repeatedly with 50 % aqueous ethanol solution and dried [32,33].
2.1. Materials Poly (L-lactide) (PLLA, Mw = 150,000) was purchased from Jinan Daigang Biomaterial Co. Ltd. (Jinan, China). 1, 6-hexanediamine, 2propanol, methylalcohol (99.8 %), acetonitrile, 2-(dimethyl-amino) ethyl methacrylate (DMAEMA, 99.8 %), triethylamine (99.8 %), tetrahydrofuran (99.8 %), hydroquinol (99.0 %), CuBr2 (99.0 %), CuBr (99.0 %), sodium hydroxide and hydrochloric were all purchased from Haihong Regent Corporation. 1, 1, 4, 7, 7Pentamethyldiethylenetriamine (PMDETA, 98 %) was purchased from J &K Chemical Ltd. (Shanghai, China). 2-Bromoisobutyryl bromide (BIBB, 98 %) was purchased from Sigma (USA). 2-Chloro-1,3,2-dioxaphospholane-2-oxide (95.0 %) was purchased from Tokyo Chemical Industry. 2-(Methacryloyloxy) ethyl Choline Phosphate (MCP) was synthesized following our previous report [21]. Ultrapure water with resistivity of 18.2 MΩ · cm was used throughout.
2.5. Characterization The chemical composition of the PLA and PLA-PMCP films were determined by X-ray photoelectron spectroscopy (XPS, AXIS Ultra DLD, Kratos, Japan) with a focused monochromatic Al Kα source (1486.6 eV) for excitation [34]. The FTIR measurements were performed using an FTIR spectrometer (ATR-FTIR, PerkinElmer, Waltham, MA, USA). The surface topography and roughness of films were observed with atomic force microscopy (AFM, MFP-3D-BIO, Oxford Instrument, UK). The static water contact angles were used to test the hydrophilicity of the surfaces of these PLA films, using the sessile drop method on a contact
2.2. Preparation of PLA films PLA particles (3.75 g) were dissolved in 75 mL of anhydrous dichloromethane and mixed uniformly and defoamed by ultrasonic. Then, the solution was poured in three glass molds and volatilized at room temperature for 24 h. The thickness of the films was approximately 2
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 2. (A) XPS spectra of pristine PLA, PMCP immobilized PLA-PMCP films. Part of the graph (binding energy ranging from 0 to 450 eV) was enlarged for details. (B) High resolution XPS N 1s spectra and curve fitting of PLA-PMCP2 and PLA-PMCP4 films and P 2p core-level spectra of the PLA-PMCP2 and PLA-PMCP4 films, respectively.
2.7. Degradation studies in vivo
angle goniometer (DSA100, Gumbo Ham bur, Germany) [35]. The molecular weights and molecular weight distributions of the films were measured by gel permeation chromatography (GPC, Waters1424). The X-ray diffraction ((XRD, Dmax 1400, 40 kV, 110 mA, Japan) was used to detect the effect of modification on the crystal structure of the films. Thermogravimetric Analysis (TGAQ50 TA Instruments, New Castle, DE, USA) was used to detect the thermal properties of the films. Thermal analysis was measured by DSC (TGAQ50 TA Instruments, New Castle, DE, USA), and the percentage crystallinity (χc) was calculated using equation:
The animal experiments were approved by the Animal Research Committee of the University [37]. The in vivo degradation test of the PLA and PLA-PMCP films were carried out on 24 male SD rats (8 weeks, 200−220 g). The all films were cut into small discs with a diameter of 1 cm and subcutaneously immersed in 75 % alcohol for 1 h, and then sterilized for 1 h by ultraviolet light. The operating table and the surgical instruments were sterilized and disinfected before use. Rats were fasted for more than 12 h before the subcutaneous implantation, then anesthetized through peritoneal injection with 10 % chloral hydrate solution at 3.5 mL/kg. After shaving and disinfection, the sterilized PLA and PLA-PMCP films were implanted into the back where two perpendicular incisions were created. The wounds were closed with silk sutures and then caged. At 2, 4 and 8 weeks, the specimens were harvested. The harvested samples were weighed after dried, finally the SEM was used to detect the morphology of the degradation in vivo.
χc= ΔHf/ΔHf0 x 100 Where χc is the degree of crystallinity, ΔHf is the heat of fusion of the sample, and ΔHf0 is the heat of fusion of 100 % crystalline polymer. The value used for ΔHf0 was 93 J/g [36].
2.6. Degradation studies in vitro 3. Results and discusssion Polymer degradation behavior in different environments were quantified by dry weight loss. PLA and PLA-PMCP films (W0) were hydrolyzed in 10 mL PBS (0.1 M, pH = 7.3 ± 0.1)、hydrochloric acid (pH = 1) and sodium hydroxide (pH = 13) solutions at 37℃ with 110 rpm in water bath, respectively. The solutions were refreshed every 24 h. At each time point, recording the mass loss of the films after rinsed with ultra-pure water and dried, followed by weighing (Wt). The mass loss was calculated according to the following equation:
3.1. Surface modification and characteristics As shown in Fig. 1, the aminolysis reaction was carried out with a mixed solution of 1, 6-hexanediamine/isopropanol, which can break the lipid bonds on the surface of PLA to form amide bonds and free hydroxyl and amino groups [31,32,38]. The films containing hydroxyl and amino groups on the surface further reacted with 2-bromoisobutyryl bromide (BIBB) to obtain a macromolecular initiator PLA-Br, then initiating an ATRP reaction to obtain a PMCP polymer brush. It is noted that the ingredient was hydroxyl terminated polylactic acid (OH-PLAOH) which can enhance the reactivity in this work.
Weight loss (%) = [(W0-Wt)/W0] ×100 % Where, W0 is the initial weight and Wt is the residual weight of the dry films at time t. After hydrolyzed in sodium hydroxide (pH = 13) solutions at 4 h and 20 h, the surface morphologies of the PLA and PLA-PMCP films were observed by SEM.
3.1.1. Composition and content of surface chemical elements In order to evaluate the chemical composition of the surface of the 3
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 3. High resolution XPS C 1s spectra and curve fitting of pristine PLA, PLA-NH2 and N 1s core-level spectra, PLA-Br and Br 3d core-level spectra, PLA-PMCP2, and PLA-PMCP4 films, respectively.
PLA films after modification at various stages, XPS was used to detecte and calculate the elements of C, O, N, P, Br. XPS wide-scan spectra detected the obvious characteristic peaks of N 1s (about 400 eV), Br 3d (68.8), and P 2p (133.6 eV) on the surface of PLA-PMCP2 and PLAPMCP4 films after ATRP reaction (Fig. 2A), indicating that the PMCP polymer brush has been successfully grafted. Fig. 2B showed the N 1s core-level spectra of PLA-PMCP2 and PLA-PMCP4, there were two distinct signals at about 399.3 and 402.1 eV, among them the latter belongs to the signal of N+ which corresponding to quaternary amine salt (N+R4) [20] of CP groups for PLA-PMCP. The P 2p signal at BE of about 133.6 eV were detected after ATRP reactions, both [N+]/[N] and [P]/ [C] were increased from 0.62 to 0.78, 5.1 × 10-3 to 8.5 × 10-3 with the amount of added monomer MCP increased, respectively. Fig. 3 showed the C 1s core-level spectra of the original PLA, PLANH2, PLA-Br, PLA-PMCP2 and PLA-PMCP4 surface successively. The C 1s core-level spectra can be curve-fitted into three peak components with binding energies at 284.6, 286.4, and 288.7 eV in general, attributable to the CeH, CeO and O]CeO, respectively. However, the peak of O]CeN at about 287.0 eV was appeared on the surface of PLANH2 and PLA-Br with the reaction proceeds that associated with the linkages between the PLA film surface and ATRP initiators. In addition, the N 1s and Br 3d core-level spectra signal at BE of about 400 and 69 eV were characteristics of covalently bonded ammonia and bromine, respectively, the [N]/[C] ratio on PLA-NH2 surface is 4.3 × 10-2 and the [Br]/[C] ratio on PLA-Br surface is 5.5 × 10-3 (determined by sensitivity factor corrected N 1s and C 1s core horizontal area ratio) are consistent with previous studies [32,33]. What’s more, the appearance of CeN peak at 285.0 eV means that the successful grafting of PMCP onto PLA surface. In this work, the grafting yield (GY) of PMCP is defined as GY = (Wa - Wb)/A, where Wa and Wb are the weights of the dry film before and after modification, respectively, and A is the film area [32]. The GY is 11.3 and 27.1 μg/cm2 of PLA-PMCP2 and PLA-PMCP4, respectively. It can also see that the P content increased from 0.32 % to 0.52 % with the increased amount of adding MCP monomer, but the Br content in the initiation sites decreased gradually as shown in Table 1, because the PMCP polymer brush becomes longer and longer while the initiation
Table 1 XPS atomic concentrations of the chemical components of pristine PLA, PLANH2, PLA-Br, PLA-PMCP2 and PLA-PMCP4 films. Samples
C
O
N
Br
P
PLA PLA-NH2 PLA-Br PLA-PMCP2 PLA-PMCP4
65.79 61.93 61.66 63.33 61.63
34.21 35.41 36.70 33.37 33.63
0 2.66 1.30 2.89 1.27
0 0 0.34 0.09 0.05
0 0 0 0.32 0.52
sites was entraped within the polymeric chains. Also, the oxygen percentage did not significantly change during the surface modification while the carbon was noticeably reduced compared to pristine PLA. In addition, the changes of functional group had been detected by ATRFTIR. As shown in Fig. S1, the peak intensity of the C]O stretching vibration at 1749 cm−1 is gradually increased with the increase of surface polymer brush. These results also indicated that the PMCP has been grafted successfully on the PLA surface by SI-ATRP.
3.1.2. Determination of amino groups The ninhydrin method was used to qualitatively and quantitatively detect the existence of NH2 groups on the aminolyzed PLA films. As presented in Fig. S2a, the absorbance after the ninhydrin reacted with 1,4-dioxane–isopropanol (1:1, v/v) solution containing 1,6-hexanediamine of known concentration in the range of 450−650 nm and the maximum absorption peak after the color reaction of the aminolyzed films with ninhydrin at 560 nm. The standard curve: y = 0.106x + 0.096, R2 = 0.9986 obtained by plotting the absorbance (y) versus amino concentration (x) in 560 nm and the absorbance of aminolyzed PLA films reacted with Ninhydrin at 560 nm was shown in Fig. S2b. Through these experiments, we determined that when the membrane area is fixed and the concentration of 1,6 hexanediamine is 3.34 × 10-2 mg/mL, the calculated amino density is 2.2 × 10-7 mol/cm2. Calculation method: ρ(NH2) = N x V/S, where N is the concentration of the amine groups on the aminolyzed PLA films, which was obtained by referring to the standard curve when the ninhydrin is colored and 4
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 4. Surface morphology of PLA and PLA- PMCP films.
dissolved in the mixed solvent, V is the volume of the color developing solution and S is the surface area of aminolyzed PLA films.
hydrophilicity of these surfaces demonstrated that the grafting reaction was successful.
3.1.3. Surface mopography and roughness Fig. 4 and Table S1 showed the statistical results of two-dimensional and three-dimensional topography and roughness of samples before and after modification. As can be seen, the surface of the initial PLA film was relatively smooth, and the surface was more roughness when the PMCP polymer was grafted. Meanwhile, as the amount of MCP monomer input increased, the molecular weight of modified films increased slightly (measured by GPC, shown in Fig. S3 and Table S2), and lead the surface roughness also increased due to the addition of PMCP brushes.
3.2. Effect of surface modification on its structure and properties
3.1.4. Water contact angles The hydrophilicity of the surface is also changed during the modification. As shown in Fig. 5, compared with the pristine PLA, the static water contact angle of PLA-NH2 decreased from about 87.0° to 68.5°. After introducing BIBB onto the PLA-NH2 surface, the water contact angle increased slightly because of the hydrophobicity of the initiator. With the surface ATRP reaction introduced onto the PLA-Br surface, the water contact angle sustainably decreased to 61.0° and 51.5° for PLAPMCP2 and PLA-PMCP4, respectively. The hydrophilicity of the surface of the PLA films were improved obviously after introduced the PMCP polymer brushes, and the more polymer brushes the higher hydrophilicity. In summary, the results of XPS and the change of
Some previous studies on the hydrolytic degradation of poly (lactic acid) suggest that PLA degrades faster in alkaline environments [39]. The catalysts act to accelerate the degradation of PLA, the alkali or lipase usually added to promote degradation. The proteinase K is the most widely used enzyme which can degrade PLA, but proteinase K can only degrade non-crystalline PLLA segments, therefore, the addition of alkali is generally used for accelerated the degradation of PLLA. As shown in Fig. 6A, the degradation studies in phosphate buffer solutions (100 mM, pH = 7.3 ± 0.1) and hydrochloric acid (pH = 1) showed slight weight loss (barely degraded) after incubating samples at 37℃ for three months, while the degradation in NaOH solutions (pH = 13) just need several days for complete degradation. The hydrolytic degradation rate of original PLA films in all conditions was higher than that of PLAPMCP2 and PLA-PMCP4 films, even though they are more hydrophilic. For instance, the pristine PLA films have degraded 73.6 % after 20 h, while PLA-PMCP2 and PLA-PMCP4 degraded 36.4 % and 23.5 % in NaOH solutions, respectively. Moreover, it has been reported that the PLA film has different degradation mechanisms in several solutions, such as the hydrolysis is mainly the surface erosion in alkaline solution and noumenon erosion in phosphate buffer and acidic solution. The durability in acidic solution is similar to neutral conditions and higher than in alkaline environments [40,41]. These results are also consistent with our previous work about PMCP modified polycaprolactone (PCL) as a tissue engineering material [21]. It is well known that for partially crystallized PLLA, water first penetrates into the amorphous region, causing the ester bond to break, and when most of the amorphous region has degraded, the edge begins to degrade toward the center of the crystalline region and it is driven via nucleophilic attack of hydroxide ions on carbonyl carbon [42,43]. Therefore, there are several possible difficulties when the modified films degraded in solutions: (1) the zwitterion first forms a hydration layer with water molecules, making
The XRD, DSC and TG had been used to detect the effect of modification on the crystal structure and thermal properties of the PLA and PLA-PMCP films. As the results showed in Fig. S4, there were no significant influence on the structure and crystallinity of the films before and after modification. 3.3. Degradation study in vitro
Fig. 5. Static water contact angles of Pristine PLA, PLA-NH2, PLA-Br, PLAPMCP2 and PLA-PMCP4 films. 5
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 6. (A) Degradation profiles of pristine PLA, PLA-PMCP2 and PLA-PMCP4 films at 37℃ in PBS (pH = 7.3), hydrochloric acid (pH = 1), NaOH solution (pH = 13) for 20 h and degraded completely and (B) SEM images of pristine PLA, PLA-PMCP2 and PLA-PMCP4 films, before and after degradation in NaOH solution for 4 h and 20 h, respectively.
6
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
Fig. 7. (A) Biodegradation profiles of PLA and PLA-PMCP films in vivo and (B) SEM images of pristine PLA, PLA-PMCP2 and PLA-PMCP4 films, implanted in vivo for 2 weeks, 4 weeks and 8 weeks, respectively.
in C content and a decrease in O content.
the water molecules less permeable; (2) there are less ester bond exposed on the film surface as they are covered by zwitterionic CP groups; (3) the slightly increased molecular weight of the PLA films after PMCP modification. As a result, the PMCP polymer brush grafted on the surface has a certain “protective” effect on the surface which weakens the nucleophilic attack of the OH- ion on the ester bond, thereby delaying the degradation rate. Furthermore, the more the grafted PMCP, the degradation rate of PLA-PMCP films are more tardiness. Scanning electron microscopy (SEM) was used to observe the surface morphology of modified PLA films before and after degradation. As shown in Fig. 6B, the PLA, PLA-PMCP2 and PLA-PMCP4 films were smooth before degradation, however, as the degradation time prolonged, a large number of gaps and pores began to appear on the surface of the PLA films, while the modified PLA- PMCP2 and PLA-PMCP4 were relatively intact. The SEM results showed that the erosion rate of PLA-PMCP4 was lower than the PLA-PMCP2 and PLA under the same conditions. Among them, PLA-PMCP4 film exhibited the best degradation resistance in terms of morphology, which also corresponds to the degradation weight loss results shown in Fig. 6A. These results further confirmed that the introduction of hydrophilic MCP onto the surface of PLA can improve its hydrophilicity while delay the degradation rate at the same time. In this study, the mechanism of resist degradation also be discussed using the chemical derivatization technique in XPS and analysed on the PLA-PMCP4 films and degradation in NaOH for 4 h and 20 h, respectively. As the C 1s core-level spectra of the films showed in Fig. S5, the peak intensity of CeO and O]CeO were decreased gradually compared with pristine PLA-PMCP4 film, which means the ester group fracture and the degradation of the films. Fig. S6 showed the N 1s corelevel spectra of the modified films after degradation in over time. The results showed that with the prolongation of degradation time, the proportion of [N+]/[N] on the surface of the modified films were decreasing, which further confirms the surface protection of PMCP and delayed its degradation. It can also see in Table S3 that as the degradation progressed, there was no significant difference between the ratio of C and O on the surface compared with pristine PLA-PMCP4 films at 4 h, but at 20 h, the more decrease of [N+]/[N] aggravated the degradation tendency of the films, and the content of C was increased obviously from 63.39 %–70.40 %, while the content of O gradually decreased from 34.67 % to 29.13 %, respectively. In summary, due to the protective effect of the PMCP polymer brush on the PLA surface, the degradation of PLA had been delayed for a period of time, when the protection of PMCP fails, the degradation of PLA will lead to an increase
3.4. Degradation studies in vivo Rats subsided the anesthesia when the pristine PLA and PLA-gPMCP films were implanted for 2 h and there was no obvious swelling or bleeding. After 24 h, they began to eat and walk without any rat death. All of them were treated with neck removal during the planned time. A representative macroscopic image of the surgical wound of the rats at different times after implantation of the films are shown in Fig. S7. At 2 weeks, the incisions healed substantially, after 4 and 8 weeks the incisions were completely healed and hardly detectable. The hair of the rats gradually grew and covered the surgical area. In addition, there was no significant difference in the macroscopic morphology of all the films before and after implantation at various times. The in vivo biodegradation profiles of pristine PLA and PLA-PMCP films were evaluated by measuring weight loss (%) over time (Fig. 7A). There were both chemical and enzymatic degradation of PLA in vivo, and the degradation rate in vivo was faster than that in vitro, which was similar to other polyester materials [44–46]. The mass loss of the films was faster in the first two weeks, and then the change was relatively flat and slightly different from in vitro degradation, which may be due to the metabolism of the body and the complex physiological environment in vivo. However, the PLA-PMCP films also exhibited slower degradation in vivo than that of PLA film, which was similar to the degradation in vitro, and that could provide a relatively stable structure at the beginning of implantation, which would be very meaningful in tissue engineering. The SEM images of the films before and after implantation in vivo have showed in Fig. 7B. We observed that the surfaces of all films are relatively smooth and there was no significant difference before implantation. As time went on, the surface of the PLA films began to crack at 2 weeks, while the morphology of the PLA-PMCP films still integrity. The cracks on the surface of PLA film had continued to expand, and tiny cracks also appeared on the PLA-PMCP films with the increase of implantation time. When at 8 weeks, the surface of PLA consisted of many small pieces, PLA-PMCP films were also increased a lot of cracks. In addition, the surface of the PLA film had the most cracks, followed by PLA-PMCP2 and PLA-PMCP4 films which indicated that the PLA film had the most degradation and the PLA-PMCP film had the best degradation resistance, which was consistent with the trend of mass loss in vivo. In summary, the PLA-PMCP films had a slower degradation rate than PLA film in vitro and in vivo which have exhibited excellent 7
Colloids and Surfaces B: Biointerfaces xxx (xxxx) xxxx
Q. Xin, et al.
properties of degradation resistance.
[16] L. Xu, K. Crawford, C.B. Gorman, Effects of temperature and pH on the degradation of poly (lactic acid) brushes, Macromolecules 44 (2011) 4777–4782. [17] P. Gunatillake, R. Mayadunne, R. Adhikari, Recent developments in biodegradable synthetic polymers, Biotechnol. Annu. Rev. 12 (2006) 301–347. [18] X. Yu, Z. Liu, J. Janzen, I. Chafeeva, S. Horte, W. Chen, R.K. Kainthan, J.N. Kizhakkedathu, D.E. Brooks, Polyvalent choline phosphate as a universal biomembrane adhesive, Nat. Mater. 11 (2012) 468–476. [19] X. Chen, T. Chen, Z. Lin, X. Li, W. Wu, J. Li, Choline phosphate functionalized surface: protein-resistant but cell-adhesive zwitterionic surface potential for tissue engineering, Chem. Commun. (Camb.) 51 (2015) 487–490. [20] Y. Jiang, Y. Su, L. Zhao, F. Meng, Q. Wang, C. Ding, J. Luo, J. Li, Preparation and antifouling properties of 2-(meth-acryloyloxy) ethyl cholinephosphate based polymers modified surface with different molecular architectures by ATRP, Colloids Surf. B 156 (2017) 87–94. [21] X. Chen, Z. Lin, Y. Feng, H. Tan, X. Xu, J. Luo, J. Li, Zwitterionic PMCP-modified polycaprolactone surface for tissue engineering: antifouling, cell adhesion promotion, and osteogenic differentiation properties, Small (2019) e1903784. [22] F. Huang, C. Ding, J. li, Resisting protein but promoting cell adhesion by choline phosphate: a comparative study with phosphorylcholine, J. Bioresour. Bioprod. 3 (2018) 3–8. [23] W. Wang, B. Wang, S. Liu, X. Shang, X. Yan, Z. Liu, X. Ma, X. Yu, Bioreducible polymer nanocarrier based on multivalent choline phosphate for enhanced cellular uptake and intracellular delivery of doxorubicin, ACS Appl. Mater. Interfaces 9 (2017) 15986–15994. [24] X. Chen, H. Shang, S. Cao, H. Tan, J. Li, A zwitterionic surface with general celladhesive and protein-resistant properties, RSC Adv. 5 (2015) 76216–76220. [25] X. Chen, M. Yang, B. Liu, Z. Li, H. Tan, J. Li, Multilayer choline phosphate molecule modified surface with enhanced cell adhesion but resistance to protein adsorption, Langmuir 33 (2017) 8295–8301. [26] Z. Wu, H. Chen, X. Liu, Y. Zhang, D. Li, H. Huang, Protein adsorption on poly(Nvinylpyrrolidone)-modified silicon surfaces prepared by surface-initiated atom transfer radical polymerization, Langmuir 5 (2009) 2900–2906. [27] G. Cheng, Z. Zhang, S. Chen, J.D. Bryers, S. Jiang, Inhibition of bacterial adhesion and biofilm formation on zwitterionic surfaces, Biomaterials 28 (2007) 4192–4199. [28] W.J. Yang, T. Cai, K.G. Neoh, E.T. Kang, G.H. Dickinson, S.L. Teo, D. Rittschof, Biomimetic anchors for antifouling and antibacterial polymer brushes on stainless steel, Langmuir 27 (2011) 7065–7076. [29] D.J. Siegwart, J.K. Oh, K. Matyjaszewski, ATRP in the design of functional materials for biomedical applications, Prog. Polym. Sci. 37 (2012) 18–37. [30] Q. Yang, J. Tian, M. Hu, Z. Xu, Construction of a comb-like glycosylated membrane surface by a combination of UV-induced graft polymerization and surface-initiated ATRP, Langmuir 23 (2007) 6684–6690. [31] Y. Zhu, C. Gao, X. Liu, T. He, J. Shen, Immobilization of Biomacromolecules onto aminolyzed poly (L-lactic acid) toward acceleration of endothelium regeneration, Tissue Eng. 10 (2004) 53–61. [32] F. Xu, Z. Wang, W. Yang, Surface functionalization of polycaprolactone films via surface-initiated atom transfer radical polymerization for covalently coupling celladhesive biomolecules, Biomaterials 31 (2010) 3139–3147. [33] F. Xu, X. Yang, C. Li, W. Yang, Functionalized polylactide film surfaces via surfaceinitiated ATRP, Macromolecules 44 (2011) 2371–2377. [34] M. Wan, H. Chen, Q. Wang, Q. Niu, P. Xu, Y. Yu, T. Zhu, C. Mao, J. Shen, Bioinspired nitric-oxide-driven nanomotor, Nat. Commun. 10 (2019) 966. [35] M. Wan, T. Xu, B. Chi, M. Wang, Y. Huang, Q. Wang, T. Li, W. Yan, H. Chen, P. Xu, C. Mao, B. Zhao, J. Shen, H. Xu, D. Shi, A safe and efficient strategy for the rapid elimination of blood lead in vivo based on a capture-fix-separate mechanism, Angew. Chem. Int. Ed. 58 (2019) 10582–10586. [36] M. Yasuniwa, K. Sakamo, Y. Ono, W. Kawahara, Melting behavior of poly (l-lactic acid): X-ray and DSC analyses of the melting process, Polymer 49 (2008) 1943–1951. [37] M. Wan, J. Zhang, Q. Wang, S. Zhan, X. Chen, C. Mao, Y. Liu, J. Shen, In situ growth of mesoporous silica with drugs on titanium surface and its biomedical applications, ACS Appl. Mater. Interfaces 9 (2017) 18609–18618. [38] Y. Gong, Y. Zhu, Y. Liu, Z. Ma, C. Gao, J. Shen, Layer-by-layer assembly of chondroitin sulfate and collagen on aminolyzed poly (L-lactic acid) porous scaffolds to enhance their chondrogenesis, Acta Biomater. 3 (2007) 677–685. [39] S.J.D. Jong, E.R. Ariasa, D.T.S. Rijkersb, C.F.V. Nostruma, J.J.K.V.D. Boschc, W.E. Hennink, New insights into the hydrolytic degradation of poly (lactic acid): participation of the alcohol terminus, Polymer 42 (2001) 2795–2802. [40] H. Tsuji, Y. Ikada, Properties and morphology of poly (L-lactide). 4. Effects of structural parameters on long-term hydrolysis of poly (L-lactide) in phosphatebuffered solution, Polym. Degrad. Stab. 67 (2000) 179–189. [41] H. Tsuji, Y. Ikada, Properties and morphology of poly (L-lactide). II. Hydrolysis in alkaline solution, Polym. Sci. Part A: Polym. Chem. 36 (1998) 59–66. [42] H. Fukuzaki, M. Yoshida, M. Asano, M. Kumakura, Synthesis of copoly (D, L-lactic acid) with relatively low molecular weight and in vitro degradation, Eur. Polym. J. 25 (1989) 1019–1026. [43] C.Y. Tham, Z.A.A. Hamid, Z.A. Ahmad, H. Ismail, Surface engineered poly (lactic acid) (PLA) microspheres by chemical treatment for drug delivery system, Key Eng. Mater. 594–595 (2013) 214–218. [44] J.M. Anderson, A. Rodriguez, D.T. Chang, Foreign body reaction to biomaterials, Semin. Immunol. 20 (2008) 86–100. [45] E.M. Christenson, J.M. Anderson, A. Hiltner, Oxidative mechanisms of poly (carbonate urethane) and poly (ether urethane) biodegradation: in vivo and in vitro correlations, J. Biomed. Mater. Res. 70 (2004) 245–255. [46] J.P. Santerre, K. Woodhouse, G. Laroche, R.S. Labow, Understanding the biodegradation of polyurethanes: from classical implants to tissue engineering materials, Biomaterials 26 (2005) 7457–7470.
4. Conclusions In the present study, different proportions of PMCP brush were successfully grafted on the PLA film surfaces via SI-ATRP. The modified films showed more resistance to degradation even though the hydrophilicity was improved obviously owing to the “protective” of PMCP polymer brush compared with pristine PLA. Moreover, with more grafted PMCP monomer, the PLA-PMCP films exhibited higher hydrophilicity and better degradation resistance. Meanwhile, the modified films in vivo also exhibited similar degradation resistance and a controlled rate of degradation in vitro environment, thereby suggesting it have a good application prospect in tissue engineering. Declaration of Competing Interest The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. Acknowledgements This work was supported by the National Natural Science Foundation of China (51803173, 51572228), Sichuan Province Science and Technology Innovation MiaoZi Funding (2018002), and Fundamental Research Funds for the Central Universities (2682019CX64). Appendix A. Supplementary data Supplementary material related to this article can be found, in the online version, at doi:https://doi.org/10.1016/j.colsurfb.2019.110630. References [1] L.S. Nair, C.T. Laurencin, Biodegradable polymers as biomaterials, Prog. Polym. Sci. 32 (2007) 762–798. [2] Y. Cheng, S. Deng, P. Chen, R. Ruan, Polylactic acid (PLA) synthesis and modifications: a review, Front. Chem. China 4 (2009) 259–264. [3] R. Kulkarni, K. Pani, C. Neuman, F. Leonard, Polylactic acid for surgical implants, Arch Surg. 93 (1966) 839–843. [4] B. Tyler, D. Gullotti, A. Mangraviti, T. Utsuki, H. Brem, Polylactic acid (PLA) controlled delivery carriers for biomedical applications, Adv. Drug Deliv. Rev. 107 (2016) 163–175. [5] N. Rotter, F. Ung, A.K. Roy, M. Vacanti, R.D. Eavey, C.A. Vacanti, L.J. Bonassar, Role for interleukin 1α in the inhibition of chondrogenesis in autologous implants using polyglycolic acid–polylactic acid scaffolds, Tissue Eng. Part A 11 (2005) 192–200. [6] T. Kissel, Y. Li, F. Unger, ABA-triblock copolymers from biodegradable polyester Ablocks and hydrophilic poly (ethylene oxide) B-blocks as a candidate for in situ forming hydrogel delivery systems for proteins, Adv. Drug Deliv. Rev. 54 (2002) 99–134. [7] M. Vert, Aliphatic polyesters: great degradable polymers that cannot do everything, Biomacromolecules 6 (2005) 538–546. [8] A.C. Albertsson, I.K. Varma, Recent developments in ring opening polymerization of lactones for biomedical applications, Biomacromolecules 4 (2003) 1466–1486. [9] E.S. Place, J.H. George, C.K. Williams, M.M. Stevens, Synthetic polymer scaffolds for tissue engineering, Chem. Soc. Rev. 38 (2009) 1139–1151. [10] M. Tang, A.F. Haider, C. Minelli, M.M. Stevens, C.K. Williams, Lactide polymerization co-initiated by carbohydrate esters and pyranoses, J. Polym. Sci. Part A: Polym. Chem. 46 (2008) 4352–4362. [11] X. Jiang, M.R. Smith, G.L. Baker, Water-soluble thermoresponsive polylactides, Macromolecules 41 (2008) 318–324. [12] C. Gottschalk, F. Wolf, H. Frey, Multi-arm star poly(L-lactide) with hyperbranched polyglycerol core, Macromol. Chem. Phys. 208 (2007) 1657–1665. [13] J. Ji, H. Zhu, J. Shen, Surface tailoring of poly (dl-lactic acid) by ligand-tethered amphiphilic polymer for promoting chondrocyte attachment andgrowth, Biomaterials 25 (2004) 1859–1867. [14] A. Hoglund, M. Hakkarainen, U. Edlund, A.C. Albertsson, Surface modification changes the degradation process and degradation product pattern of polylactide, Langmuir 26 (2010) 378–383. [15] A.M. Reed, D.K. Gilding, Biodegradable polymers for use in surgery poly (glycolic) / poly (lactic acid) homo and copolymers: 2. In vitro degradation, Polymer 22 (1980) 494–499.
8