HOW TO MAKE POLYESTERS MORE PRONE TO CELL ADHESION Björn Atthoff and Jöns Hilborn Department of Material Chemistry Uppsala University Box 538 75121 Uppsala, Sweden Introduction The interface between synthetic and living matter is a very active research field for scaffolding materials in tissue engineering and regenerative medicine1. Interaction between cells and materials, the biomechanics and the mechano transduction, are important factors for appropriate integration of biodegradable polymers2. Scaffolds provide temporary mechanical support, hence should closely match mechanical properties of tissue. Soft tissue repairs require a soft compliant material3, while stiffer scaffolds should be the proper choice for cartilage or bone repair. Regardless of the application, adhesion between cells and scaffold are required for cell seeding, cell spreading, and to transmit loads between the scaffold and the surrounding tissue. Extra Cellular Matrix (ECM) components such as collagen, laminin and fibronectin have been widely used to prepare scaffolds, since these materials are excellent for cell adhesion. However, their drawbacks are the lack of adequate mechanical properties and inability to be readily processed into stable three-dimensional forms. Scaffolds made from biodegradable polymers which are acceptable for in vivo use, e.g. poly(lactic acid) (PLA), poly(lacticco glycolic-acid) (PLGA), and poly(glycolic acid) (PGA), have the requisite mechanical properties and can be processed into various shapes and forms. However, polymer surfaces do not readily interact with cells, which limit their effectiveness as scaffolding materials. Since scaffold surfaces should allow cell adhesion and development, attempts are made to mimic natural structure and properties, so that cells will attach to and proliferate on the scaffold, often, this is accomplished by immobilizing specific proteins or peptide sequences onto the polymer. One approach to surface modification is graft polymerization with an active monomer, like acrylic acid4. Other activation techniques that have been reported are exposure to radio frequency glow discharge, (plasma)5 and surface modification with PEO polymers6. For protein adhesion to occur, the proteins must first adsorb onto the polymer surface. Currently, the Andrade & Hlady model is one of the most accepted models for protein adsorption7, where the initial weak interaction is followed by protein denaturation promoting stronger adhesion. The mechanisms governing the protein to polymer interactions involve polyelectrolyte absorption8, hydrophobic9 and ionic interactions10, depending on the substrate and type of proteins. PGA, PLGA and PLA belong to a class of biodegradable polyesters whose surfaces may be altered by hydrolysis, alternatively, aminolysis may be utilized for surface functionalization. The effects of temp, pH and conc. of the substances used, were studied to optimize the conditions for cell adhesion and cell growth. Data from literature points to that unmodified biodegradable polyesters are apt for cell adhesion and cell growth11. This study have examined whether there is a need to modify the surfaces of polyesters. PET, PLA and PGA have been studied for good protein adsorption by utilizing some of the methods discussed above. This study demonstrates that surface modifications are not necessary for attachment of proteins to bio-degradable polyesters, while kinetics may be slow, in time the proteins will properly adhere. However, a surface treatment improves proper protein adhesion for non-degradable polyesters. Experimental Materials. PET (Reliance Industries Ltd.), bovine collagen type 1 (Symatese Biomatériaux), phenol (>99.5%), acetic acid (100%), stannous 2ethylhexanoate (~95%), phenethyl alcohol (99%), sodium hydroxide (97+%), 1,2-dichlorethane (99%) and (3-mercaptopropyl)trimethoxysilane (95%) (Aldrich), L-lactide S and glycolide (Boehringer Ingelheim), glutaraldehyde 50% and (3-aminopropyl) trimethoxysilane (97%) (Lancaster Synthesis Ltd.), Tissucol, aprotinin, thrombin, tissucol- and thrombin-buffer (Baxter AG). Instrumentation. JEOL ECP 400MHz NMR, LEO 1550 FEG SEM, Philips XL30 ESEM FEG, Ubbelohde capillary viscometer 0c from Schott, Physical Electronics ESCA Quantum 2000, SPI-DRYTM critical point dryer, POLARON SC7640 sputter coater, Quartz crystal microbalance QCM D300 (Q-Sense, Sweden) with quartz crystals: QSX 303 - SiO2 and QSX 301Standard Gold, Compression-molding machine (Rudolf Westerberg AB, Sweden), KW-4A spin coater (SPI Supplies), Unilab glove box (Mbraun) and an in-house constructed and built polymerization reactor. Polymer synthesis. PGA: Glycolide, 50g (0.43 mol), phenethyl alcohol, 60.44mg (0.495mmol) and Sn(Oct)2, 10mg (25µmol) and a magnetic stir bar was charged dry into a 100ml Erlenmeyer flask and polymerized at 170°C. When the viscosity increase stopped the magnetic stirrer, the heat was switched off and flask was allowed to cool to room temperature (RT). The solid polymer was light brown having an I.V. of 3.5. PLA: L-Lactide, 1000g (6.9mol) and Sn(Oct)2, 0.5g (1.2mmol) were charged dry into the polymerization reactor and polymerized at 120°C for 72 hours. The solid polymer was white, having an I.V. of 7.38. Disc preparation and protein immobilization. PLA and PGA discs; 1g of polymer, was compression molded at 200°C (PLA) and 220°C (PGA) at 60kN over 100cm2 for 60s, giving discs with a thickness of 0.3 mm. PET discs; 0.2g of PET dissolved in 4g of phenol, 50 °C, and diluted with 4g of 1,2-dichlorethane. The solution was casted onto glass under nitrogen, 90°C, 1h, giving discs with a thickness of ~0.2mm. Surface hydrolysis, discs were submerged in to 2.5 M NaOH, 50°C, 30min (PLA), 5min (PGA) and 2h (PET), followed by extensive rinsing in deionized water. Collagen: The discs were treated in 3mg/mL collagen solution, RT, 18h, before being placed in PBS, pH 7.4, 1h. The discs were washed in deionized water two times and then dried for ESCA or fixated using 5% glutaraldehyde in PBS, dehydrated in ethanol, super critically extracted, coated with 90% Au and 10% Pd and then examined using SEM. Fibrin: Lyophilized tissucol, 0.3g, was dissolved in 2mL of aprotinin solution, diluted with tissucol buffer to a concentration of 10mg/mL. The discs were submerged into that solution, RT, 18h. A thrombin solution was prepared by dissolving lyophilized human thrombin, 500IE, in calcium chloride, 1ml, and the solution was diluted to 10IE/ml by the addition of thrombin dilution buffer. The discs were placed into the thrombin solution for 1 hour, followed by washing in PBS. Samples were then washed and treated the same way as the collagen samples for analysis. Quartz Crystal Microbalance experiment. The crystals were treated in an ethanol solution of 5mM (3-mercaptopropyl)trimethoxy-silane, 40°C, 30min, followed by washing in ethanol. The crystals were then submerged in 2% (3-aminopropyl)trimethoxysilane in 95/5 ethanol/water solution, with pH 4.5 – 5.5, 5min, followed by washing in ethanol and drying under Ar. PGA and PLA were dissolved in chloroform, 5wt%, PET in (50/50) phenol/1.2dichlorethane, 2.5wt%. The polymer solutions was spin-coated, 3000rpm, 20s, onto the crystals, followed by heating under Ar, 10min, 130°C (PGA and PLA), 220°C (PET). One at a time the crystals were used in the instrument. After 1h in deionized water, the protein solutions, 0.5 mg/mL (collagen) and 10 mg/mL (fibrinogen), were injected. The absorption was measured as frequency decrease with time. After stabilization, a deionized water and 1N acetic acid wash was preformed. Results and Discussion Protein adsorption depending on surface modification, of three different polyesters has been evaluated, namely: PET, PLA and PGA. Under physiologic conditions, pH 7.4, PET is non degradable, PLA degrades slowly, while PGA rapidly degrades. The reason for the differences in degradation behavior of the three polymers, originate from the stability of the ester group. Generally aromatic esters are less reactive than aliphatic ones. PET has an aromatic-aliphatic ester while PLA and PGA are aliphatic. The difference in rate of hydrolysis between PGA and PLA is attributed to the methyl group, making PLA less reactive due to steric hindrance. The reactivity difference has a direct influence on the protein adhesion rate. The proteins used in this study are Collagen type 1 and fibrinogen, which in combination with thrombin gives fibrin. Collagen is a high molecular weight polymer that is normally kept in solution at low pH. When the pH is raised to 7.4, at room temperature, the collagen forms a gel, as the molecules arrange themselves in a fibrous network. A single collagen molecule has a diameter of 1.5 nm. These molecules arrange into triple helices that forms fibrils which in turn build up the collagen fibers, consisting of around 270 collagen molecules giving a fiber diameter of around 40nm. Polymer Preprints 2005, 46(1), 405 Fibrinogen, that stays in solution at physiological, instead, fibrinogen polymerizes with the help of thrombin to form a cross-linked fibrin network. One option to from stronger links between the polyester and the protein than is capable from hydrophobic interactions alone is to hydrolyze the polyester which produces functional alcohols and carboxylic acids. These functional groups allow additional interactions between the polymer surface and the protein. Specifically, ionic interactions between the negatively charged carboxylate anion on the surface and the positively charged ammonium cations found in the proteins are believed to be important. The surface modification used in this study is basic hydrolysis, where the polyesters establishes carboxylates and hydroxyl group on the surface of the polymer, to which the proteins could attach by polyelectrolyte adsorption. Aminolysis from the protein on the ester may also take place for the degradable polyesters. Since ester group reactivity of the different polymers varies considerably, different hydrolysis times were used in this study, PGA, 5min, PLA, 30min, and PET, 2h. The different times were chosen on the basis of a screening degradation study. After the protein immobilization and washing, the presence of proteins on the samples surfaces were confirmed using ESCA and the morphology of the protein was studied by SEM, while the adsorption kinetics were determined using quartz crystal microbalance (QCM). ESCA result shows that all hydrolyzed polymers, adsorbed a significant amount of protein, indicated by a nitrogen content of ~15%, Table 1a. Interestingly, for the non-hydrolyzed samples a difference can be seen between degradable and non-degradable polymers. PET shows nitrogen content of 6%, while PLA and PGA both show ~15%, which indicates that more protein has been immobilized on the degradable polymers. The same pattern was observed for fibrinogen immobilization, Table 1b. Table 1a. ESCA, Collagen Table 1b. ESCA, Fibrinogen PET Non-hydrolyzed (%N) 6.0±2.8 Hydrolyzed (%N) 13.7±1.6 PET Non-hydrolyzed (%N) 7.8±1.7 Hydrolyzed (%N) 12.6±0.9 PLA 13.5±1.8 15.3±0.9 PLA 14.1±2.8 14.8±1.9 PGA 14.7±0.5 13.7±2.4 PGA 15.6±0.2 14.5±1.4 These results were verified with SEM, that showed that non hydrolyzed PET not had adsorbed as much protein as the hydrolyzed. Protein structures similar to those seen on hydrolyzed PET are observed both on the hydrolyzed and the non-hydrolyzed PLA and PGA. It seems likely that surfaces having charged species would attract proteins at a different rate than the purely hydrophobic surfaces. Therefore, the kinetics of immobilization was monitored using QCM. The resonant frequency of the crystal depends on the oscillating mass. Adsorbed material on the crystal increases mass and decreases frequency. The decrease in frequency is proportional to the materials mass, so the QCM operates as a very sensitive balance, allowing the monitoring of protein adsorption in real time. From protein adsorption curves it is evident that non hydrolyzed PET adsorbs less protein, and at slower rate, than hydrolyzed PET, which is why the hydrolyzed curve in figure 3a has a larger frequency drop than the non hydrolyzed one. The adsorption behavior of PLA in figure 3b, is quite different. Non hydrolyzed PLA adsorbs protein at a slower rate, although the difference in adsorption behavior between nonhydrolyzed and hydrolyzed PLA is less than for PET. The PLA samples should adsorb the same amount of proteins since the non hydrolyzed sample slowly will hydrolyze in the protein solution and more active binding sites are created as the polymer degrades. On the hydrolyzed sample all those sites are already there when the sample is submerged into the protein solution the immobilization is faster, but they should end up with the same amount of protein in the end. This behavior is similar for PGA although the adsorption rates are much higher, and the difference between the two curves smaller. The adsorption behavior of proteins on PGA however is difficult to measure for extended times due to its rapid degradation, almost immediately after the sample is placed in the protein solution, PGA starts to degrade at high rate. The QCM results show that the rates of immobilization varied between the hydrolysis and non-hydrolyzed samples, with a larger spread for PLA than for PGA. PET does not hydrolyze upon exposure to the protein solution and therefore a separate treatment must be employed to create an adequate number of binding sites for protein adsorption. Adsorption is controlled by polyelectrolyte adsorption, where the negatively charged carboxylate ions are attracted to the positively charged ammonium ions of the protein, as well as hydrophobic interaction. One should not exclude the possibility that aminolysis of the ester of PLA and PGA may take place to covalently link the protein to the polymer. Figure 3a. QCM, PET, collagen Figure 3b. QCM, PLA, Collagen The rate and ease of protein immobilization is directly related to the reactivity of the different esters, where PET is least reactive, followed by PLA and finally the most reactive PGA. The order of reactivity is reflected in the QCM studies, where the adsorptions for the more reactive of the two degradable esters are faster than the less reactive one. PET cannot be compared since it does not degrade in the protein solution to form more active sites, hence there is no or a very slow reactivity in the PET ester groups. Conclusion ESCA, SEM and QCM have complemented each other to give the results that suggest that surface activation is not necessary to achieve good protein adsorption onto degradable polyesters, the spontaneous hydrolysis by the ester group in these polymers are adequate enough to get the proteins to adsorb. However, a minor surface hydrolysis may increase the adsorption kinetics. For the non-degradable polyesters, there are no evident degradation in physiological, or even slight acidic environments, hydrolysis will help to break the esters, thereby increasing the number of active sights for the proteins to attach. With the help from this study, we now have methods to attach different protein to scaffold materials in order to achieve good cell adhesion and proliferation, both by providing the correct surface chemistry, but also to achieve better mechanical matching between scaffold and cells. The protein layers may act as a soft layer so that the cells might experience even a hard polymer surface as soft due to the adsorbed protein layer. References (1) Castner, D. G.; Ratner, B. D. Biomedical surface science: Foundations to frontiers, Surface Science, 2002, 500, 28-60 (2) Tomasek, J. J.; Gabbiani, G.; Hinz, B.; Chaponnier, C.; Brown, R. A. Myofibroblasts and mechanoregulation of connective tissue remodelling, Nature Reviews, 2002, 3, 349-363 (3) Oberpenning, F.; Meng, J.; Yoo, J. J.; Atala, A. De novo reconstitution of a functional mammalian urinary bladder by tissue engineering, Nature Biotechnology, 1999, 17, 149-155 (4) Gupta, B.; Plummer, C.; Bisson, I.; Frey, P.; Hilborn, J. Plasma-induced graft polymerization of acrylic acid onto poly(ethylene terephtalate) films: characterization and human smooth muscle cell growth on grafted films Biomaterials, 2002, 23, 863-871 (5) Klomp, A. J. A.; Engbers, G. H. M.; Mol, J.; Terlingen, J. G. A.; Feijen, J. Adsorption of proteins from plasma at polyester non-wovens Biomaterials, 1999, 20, 1203-1211 (6) Elbert, D. L.; Hubbel, J. A. Surface Treatments of Polymer for Biocompatibility Annual Reviews, 1996, 26, 365-394 (7) Andrade, J.; Hlady, V. Protein Adsorption and Materials Biocompatibility: A Tutorial review and Suggested Hypoteheses, Advances in Polymer Science, 1986, 79, 12 (8) Richert, L.; Boulmedais, F.; Lavalle, P.; Mutterer, J.; Ferreux, E.; Decher, G.; Schaaf, P.; Voegel, J-C.; Picart, C. Improvment of Stability and Cell Adhesion Properties of Polyelectrolyte Multilayer Films by Chemical Cross-Linking Biomacromolecules, 2004, 5, 284-294 (9) Wolde, P. R. Hydrophobic interactions: an overview, Journal of Physics: Condensed Matter, 2002, 14, 9445-9460 (10) Kato, K.; Sano, S.; Ikada, Y. Protein adsorption onto ionic surfaces Colloids and Surfaces B: Biointerfaces 1995, 4, 221-230 (11) Ishaug-Riley, S. L.; Okun, L. E.; Prado, G.; Applegate, M. A.; Ratcliffe, A. Human articular chondrocyte adhesion and proliferation on synthetic biodegradable polymer films Biomaterials, 1999, 20, 2245-2256 Polymer Preprints 2005, 46(1), 406
© Copyright 2024