HOW TO MAKE POLYESTERS MORE PRONE TO CELL ADHESION Introduction

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.
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