Correlation between heparin release and polymerization degree of organically modified silica xerogels from 3-methacryloxypropylpolysilsesquioxane

Acta Biomaterialia 3 (2007) 829–837 www.elsevier.com/locate/actabiomat

Correlation between heparin release and polymerization degree of organically modified silica xerogels from 3-methacryloxypropylpolysilsesquioxane David Tebbe, Roger Thull, Uwe Gbureck

*

Department for Functional Materials in Medicine and Dentistry, University of Wu¨rzburg, Pleicherwall 2, D-97070 Wu¨rzburg, Germany Received 20 December 2006; received in revised form 2 April 2007; accepted 2 May 2007 Available online 22 June 2007

Abstract This work aimed to investigate the use of an organically modified porous silica matrix (poly(methacryloxypropyl)–poly(silsesquioxane); P-MA–PS) as a release system for heparin. The matrices were obtained from methacryloxypropyltrimethoxysilane (MAS) via the sol–gel process under acidic conditions following photochemical polymerization and cross-linking of the organic matrix. Modulation of the polymerization degree of the organic matrix in the range 0–71% allowed adjusting the release kinetics of heparin according to therapeutic needs. It was demonstrated that higher drug loads and a decreasing polymerization degree resulted in a faster release profile of heparin, which followed a square root of time kinetic according to the Higuchi model. The hydrolytic degradation of hybrid xerogel was found to follow a zero-order kinetic whereas the heparin concentration did not show an influence on the degradation rate of the matrix. Since the released heparin retained its biological activity, the P-MA–PS matrices are of clinically interest, e.g. as coating on drug eluting coronary stents. Ó 2007 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. Keywords: Heparin; Drug release; Silica xerogels; Polymerization

1. Introduction The design and development of implantable drug-delivery systems has become increasingly attractive due to their advantages of safety, efficiency and patient convenience. A common example is the use of antibiotic-loaded polymethylmethacrylate (PMMA) bone cement for the prevention and treatment of post-operative bone infections like osteomyelitis [1,2]. Another field of application represents the coating of metallic implants with biological active molecules to improve their blood compatibility. For instance, stent surfaces were modified with drugs like heparin [3] and bivalirudin [4] to prevent side effects like blood coagulation and neointimal proliferation after implantation [5]. *

Corresponding author. Tel.: +49 931 201 73550; fax: +49 931 201 73500. E-mail address: [email protected] (U. Gbureck).

Surfaces can be modified either by covalent bonding of the drug to the metal by a silane spacer [6–8] or by embedding the active agent into a polymer matrix for the controlled release over a certain period of time [9,10]. The covalent attachment has the disadvantage that the potency of the drug can be reduced if biological active groups are sterically hindered by the immobilization [11,12]. In contrast, the efficiency of drug release from polymer matrices like PMMA polymers [13], as well as degradable polylactic acid/polyglycolic acid (PLA/PGA) polymers [14], has been proved. However, the low adhesion of the organic polymer matrix on the metallic surface [15] is unfavourable and may result in partial or full delamination of the polymer coating in the aqueous environment in vivo due to a hydrolytic cleavage of the adhesive metal–polymer interface. This effect and simultaneous degradation of the coating holds the risk of detaching bigger fragments which can cause restenosis inside blood vessels [16]. To improve the adhesion

1742-7061/$ - see front matter Ó 2007 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2007.05.002

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of the polymer matrix, bifunctional precursor molecules (Fig. 1) can be applied to form a covalent and mostly hydrolytic stable bond to a metallic surface bearing OH-groups, e.g. titanium with its native oxide film. This method is well known from dental metal–polymer joints [17,18] and was further investigated for improving the adhesion of PMMA cements to metallic prostheses [19]. Methacrylate functionalized silanes, known as adhesive agents, can also be taken to generate bulk materials by means of the sol–gel technology [20]. The resulting organic–inorganic hybrid polymers also show good adhesion on metals [21] and, moreover, have better mechanical properties compared with pure polymers. By varying process parameters (hydrolysis and polymerization conditions), the structure (density, nanoporosity) can be adjusted over a broad range [22]. Simple inorganic bulk materials consisting of nanoporous silica were already established as delivery systems for the drug heparin [23,24]. This study aimed to investigate the use of sol–gelderived organic–inorganic PMMA polymers as delivery systems for the controlled release of heparin as a model drug and adjustment of critical parameters which influence the drug release kinetics without considering the adhesion between the precursor and titanium, which has been the topic of numerous works in the past showing a high stability of the interface [17–19]. We hypothesized that the variation of the cross-linkage of the organic methacrylic groups can be used to control the release kinetics of the drug. Therefore, the precursor methacryloxypropyltrimethoxysilane was first hydrolyzed by the sol–gel process and then loaded with heparin following photochemical polymerization of the organic moieties. The amount of released heparin and its biological activity were measured photometrically by the toluidine blue test and the chromogenic substrate (Chromozym TH) method, and then correlated with the porosity and degree of polymerization of the organic matrix. 2. Materials and methods 2.1. Preparation of organic–inorganic hybrid gel films (poly(methacryloxypropyl)–poly(silsesquioxane)) The heparin containing poly(methacryloxypropyl)– poly(silsesquioxane) (P-MA–PS) samples were prepared

by a two-step synthesis starting from the precursor methacryloxypropyltrimethoxysilane (MAS). First, the trimethoxysilyl moiety was cross-linked by means of the sol–gel process, followed by photoinitiated radical polymerization of the methacrylate functionality as the second step. The sol–gel process was carried out under a dry nitrogen atmosphere with HCl as catalyst [20]. The following reagents were used: MAS (A 174; Merck), methanol (Merck), deionized water (bidistilled), hydrochloric acid (10 M; Merck), benzoinmethylether (96%; Sigma–Aldrich) and heparin sodium salt (Sigma–Aldrich). Heparin is derived from animal resources (porcine) and the average molecular weight is between 13 and 18 kDa. The biological activity of the used heparin was 173 IU mg1 as given by the distributor and validated by a chromogenic method. The molar ratio of the starting sol MAS:CH3OH:HCl:H2O ratio was 1:2:0.1:0.6. After hydrolysis of MAS with H2O and the catalyst (2 h boiling under reflux conditions) a colourless homogeneous solution was obtained, to which 0.45 g of heparin and 2.00 mmol of benzoinmethylether were then added. The concentration of heparin in the sol was varied in the range 1.7–5.0 wt.%, corresponding to 2.7–8.0 wt.% in the air-dried xerogel. The sol was cast into plastic molds (diameter 5 cm) and kept at room temperature for polycondensation, ageing and drying for 3 days. For cross-linking of the organic matrix, the drug-loaded xerogel samples were exposed to ultraviolet (UV) light (150 W, XBO Xenon lamp, Osram, light spectrum 200– 1100 nm) for 1–15 min. The successful formation of siloxane bonds by means of the sol–gel process was confirmed by infrared (IR) spectroscopy and the degree of the photochemical polymerization was measured by means of Raman spectroscopy. 2.2. Determination of the polymerization degree To determine the degree of polymerization of the P-MA– PS samples, Raman spectra of the photochemically cured xerogels were recorded and related to those of the nonpolymerized substances. Briefly, the areas of the peaks of the methacrylic C@C bond at 1635 cm1 were taken and referenced to those of the respective aromatic C@C bond at 1605 cm1. The areas of the peaks were integrated by fitting a Gaussian function. The polymerization degree of the double bonds was calculated after Eq. (1):

Fig. 1. Preparation of organic/inorganic P-MA–PS drug depots. Heparin was incorporated into the sol–gel reaction mixture after hydrolysis and partial condensation of MAS.

D. Tebbe et al. / Acta Biomaterialia 3 (2007) 829–837

DegreePolymerization ¼

  C@CPolymer 1  100 ½% C@CRaw material

ð1Þ

where C@CRaw material and C@CPolymer are the number of double bonds of raw material and of polymer, respectively. The Gaussian-fitted areas of the methacrylic and aromatic double bonds (AMethacryl, resp. AAromat) of the raw material and polymeric compounds (ARaw material, respectively APolymer) were taken to calculate the polymerization degree (Eqs. (2)–(4)). C@CRaw material ¼ C@CPolymer ¼

material ARaw Methacryl material ARaw Aromat

APolymer Methacryl APolymer Aromat

DegreePolymerization ¼

APolymer Methacryl  AAromat

Raw material APolymer Aromat  AMethacryl

ð3Þ

The total amount of dissolved heparin was measured by the colorimetric toluidine blue method [26,27]. The biological potency of heparin after dissolution from the hybrid matrixes was determined by the chromogenic substrate Chromozym THÒ [28,29]. Determination of the anticoagulant activity of heparin was done from the same dissolution samples as the quantification of the released heparin. Different aliquots of the liquid dissolution medium were used to be within the measurement range of the method and different heparin loads of the release matrices were taken to identify the influence of drug load on the biological activity. Small differences in measurement time are a result of the different starting times of the experiments. Since the variation in time is within ±10%, this has only a marginal effect on the results. Therefore, the concentrations of heparin in both the calibration solutions and samples were diluted to a maximum concentration of 0.3 U ml1, estimating a maximum activity of 173 U mg1. The calibration for this method using the chromogenic substrate was carried out according to a procedure described in literature [29]. To determine the potency of heparin dissoluted from the differently polymerized (0%, 36% and 71%) and drugloaded (50, 100 and 150 mg) matrices, aliquots of 100 ll from each of the polymerized samples with a drug load of 50 mg were taken 72 h after starting the dissolution experiment and diluted with 0.15 M NaCl solution to a total volume of 1.11 ml. Aliquots of 100 ll from the samples with a drug load of 100 mg were taken 76 h after starting the dissolution test and diluted to a total volume of 2.22 ml, respectively. Finally, aliquots of 100 ll from the samples with a drug load of 150 mg were also taken 76 h after the beginning of the experiment and diluted to a total volume of 3.33 ml. From the nine different aliquots thus obtained, a quantity of 100 ll of each was taken, mixed with 100 ll of AT-III solutin (1 U ml1) and 800 ml of Tris buffer solution (50 mM), and then poured into a glass tube. The glass tubes were pre-warmed in a water bath at 37 °C for 2–6 min. Stage I was started by blowing 500 ll of thrombin (8 U ml1) into the tube. After 30 s, stage II was initiated by adding 1.00 ml of ChromozymTHÒ substrate solution (0.75 mM) into the reaction mixture. After 60 s (90 s after the start of stage I), amidolysis was stopped

 100 ½%

2.3. Release of heparin in vitro The dissolution tests of heparin from the different hybrid matrices were carried out in a shaking water bath at 37 °C (60 shakes per min). The average weight of the samples was 2 g (diameter 5 cm). Phosphate-buffered saline (PBS) was used as dissolution medium [25], because the composition of inorganic ions as well as the pH value are close to that of human blood plasma. PBS (pH 7.4) was freshly prepared by dissolving 8.00 g (137 mmol) of NaCl, 0.20 g (2.68 mmol) of KCl, 1.11 g (9.10 mmol) of Na2HPO4 and 0.20 g (1.47 mmol) of KH2PO4 in 1 l of deionized H2O. Heparin loaded samples were immersed in 250 ml of PBS in a glass bottle and 0.33–1.00 ml aliquots of sample liquid were removed at scheduled times for heparin quantification and replaced with fresh PBS. Different aliquots were used such that the heparin concentration was laying within the measurement range of the method. Three parallel samples were examined, each data point being the mean of three values. 2.4. Porosity of the matrix The test samples were prepared by casting 0.50 ml of the sol–gel reaction mixture into plastic molds (diameter = 6.10 mm). Afterwards, the matrices were dried at room temperature for 24 h and then at 60 °C for another 24 h. Polymerization of the bulk materials was also realized as described previously. Finally, the samples were weighted dry and then immersed in 10 ml ethanol for 24 h and weighed directly after removal (to give Dm). The porosity was calculated according to Eq. (5): V Ethanol Dm  100 ½% with V Ethanol ¼ qEthanol V Probe

Degradation of the P-MA–PS xerogel matrix during the dissolution test was evaluated by measuring the amount of hydrolyzed MAS in the PBS release medium. MAS could be quantified photometrically because of its carbon–carbon double bond at 220 nm by means of UV–visible spectroscopy. Determination of dissolved MAS was done from the same dissolution samples as the quantification of released heparin. 2.6. Quantification of heparin and determination of the biological activity

ð4Þ

Porosity ¼

2.5. Degradation of the xerogel matrix

ð2Þ

! Raw material

1

831

ð5Þ

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by adding 1.50 ml of concentrated acetic acid into the mixture, for which complete mixing was essential. The absorbance was read in a spectrophotometer at 405 nm against distilled water.

1080 cm1, in conjunction with an increase of m(Si–OH) at 3700–3200 cm1 and m(Si–O–Si) at 1130 cm1 [20,21], which indicates the hydrolytic polycondensation of MAS to afford MA–PS gel films cross-linked by siloxane bonds.

3. Results and discussion

3.2. UV polymerization of the organic matrix

3.1. The sol–gel process

The radical polymerization of MA–PS was activated photochemically using UV irradiation and benzoinmethylether as initiator to afford P-MA–PS hybrid networks. The polymerization process was confirmed by Raman spectroscopy by the decrease of the vinylic C@C peak at 1635 cm1. The polymerization degree of the matrix was quantified by comparing the number of non-polymerised free carbon–carbon double bonds (C@C) in the raw material and the polymerized product. The amount of non-polymerized C@C bonds can be measured by means of 13C nuclear magnetic resonance [32], FT-IR spectroscopy [33,34] and Raman spectroscopy, whereas Raman spectroscopy has the advantage of a much higher intensity of the C@C peak compared with FT-IR spectroscopy. Within the Raman effect, the intensity of the vibration can be considered as proportional to the concentration of the corresponding bond. Since a calibration of this method cannot be performed on the basis of defined polymeric compounds, the peak of another vibration (with no variation of intensity during polymerization) of the samples has to be taken as internal standard. For this, the aromatic C@C valence vibration of the UV initiator benzoinmethylether at 1605 cm1 was suitable; in the case of polymerization without the initiator, the vibration of the ester group from the precursor molecule at 2900 cm1 was taken instead. The influence of different UV light exposure regimes on the polymerization degree with and without benzoinmethylether is shown in Fig. 3. It seems obvious

Sol–gel technology offers new possibilities for incorporating biologically active agents within bulk and coating materials at room temperature with control of their release kinetics from the matrix. The sol–gel technique is inexpensive, versatile and simple, and the xerogels generated are non-toxic and biocompatible materials [30,31]. Physical (density, porosity) and chemical properties (cross-linkage) of silica xerogels prepared by hydrolysis and condensation of the inorganic alkoxysilane moiety can be adjusted over a broad range through variation of reactants, pH and hydrolysis conditions, as well as the drying conditions of the silica sol. Previous studies have shown the effects of alternating catalysts and the amounts of water and solvent on the chemical and physical properties of P-MA–PS hybrid matrices [20,21], but no effort has been made to investigate these materials as drug-release systems. The hydrolytic polycondensation of MAS can be carried out acid- or base-catalyzed [20]. Bases like NaOH with a strong nucleophilic character have the disadvantage that they may hydrolyze the methacrylate ester, which results in cleavage and consequently the destruction of the MAS precursor molecule. For this reason, the polysilsesquioxanes in this work were processed exclusively with hydrochloric acid. The successful formation of the polysilsesquioxane network was shown qualitatively by means of Fourier transform (FT)IR spectroscopy (Fig. 2). The FT-IR spectrum of MA–PS gel film shows a decrease in the absorption peak intensity of m(–OCH3) at 2850 cm1 and accordingly m(Si–OCH3) at

Fig. 2. FT-IR spectra of (a) MAS precursor and (b) MA–PS xerogel.

Fig. 3. Polymerization degree of P-MA–PS samples dependent on the UV light exposure (with and without benzoinmethylether for UV activation). Each data point is the mean of three samples and error bars are standard deviations.

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833

Table 1 Determined porosities of the differently polymerized matrices and gradients kH with standard deviations (n = 3) from linear regression of the Higuchi plot (Fig. 6) Degree of polymerization (%)

0

36

71

Porosity

63.4 ± 10.1

42.0 ± 15.2

30.6 ± 10.7

Drug load (mg) kH Standard deviation

50 3.48 0.10

100 5.34 0.35

150 8.05 0.44

that initiation with benzoinmethylether is necessary to obtain high polymerization degrees in appropriate times. 3.3. Porosity The porosity of the differently polymerized (0%, 36%, 71%) P-MA–PS samples was determined by the absorption of ethanol (q = 0.79 g cm3 at 20 °C). The difference in weight thus caused, and thus the deducible volume of the adsorbed alcohol, was divided by the volume of the respec-

50 2.83 0.07

100 4.42 0.24

150 6.17 0.41

50 2.36 0.06

100 3.63 0.22

150 5.02 0.17

tive probes to give the corresponding porosity [35]. The results (Table 1) show that non-polymerized MA–PS samples have the highest porosity, at 63 ± 10%. Compared with pure silica (SiO2) xerogels, with porosities of approximately 50% [35], the value obtained here is higher, probably caused by the bulky organic part of the MAS precursor molecule. This methacrylic moiety leads to a less dense package of MA–PS and therefore a higher porosity. Furthermore, increasing degrees of polymerization (36%, 71%) of P-MA–PS samples are linked to strongly decreasing

Fig. 4. Release of heparin from hybrid matrixes with different polymerization degrees (0%, 36% and 71%) with a drug load of (a) 50, (b) 100 and (c) 150 mg. Each data point is the means of three samples and error bars are standard deviations.

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Fig. 5. Correlation between porosity and heparin release of hybrid gels with different drug loads after 144 h. Each data point is the mean of three samples and error bars are standard deviations.

porosities (42 ± 15% and 31 ± 11%) of the matrices. This phenomenon is caused by the additional organic cross-linkage of the methacrylic functionality in two ways. First, polymerization is associated with a shrinking process, which is a well-known characteristic of organic methacrylic acid-based polymers [36] and which was found to be 10 vol.% in our study due to dimensional changes of the samples. Secondly, polymerization of the organic moieties might have decreased the diameter of the pores inside the matrix such that ethanol no longer entered into the smallest pores due to the capillary forces. Hence, it follows that the obtained porosity values are probably lower that the actual porosity of the matrices. More precise measurements may be achieved by measuring the strut densities of the solid products using helium pycnometry. 3.4. Heparin release The release of heparin from all matrices with different drug loads and varying polymerization degrees can be considered as first-order kinetic during the dissolution period due to the finite nature of the drug depot (Fig. 4a–c). The correlation between the polymerization degree and the respective porosities with the release of heparin from the differently drug-loaded samples is shown in Fig. 5. The initial burst release from matrices is a well-known problem which is more evident for systems with higher surface area, and increases when the molecular weight of heparin decreases and the amount of embedded heparin increases [24]. As the embedded heparin is a fraction of different weights (13–18 kDa), in the first dissolution period smaller drug molecules are eluted faster than bigger ones. Furthermore, the concentration of the drug close to the matrix surface is still high, which shortens the discharge

path of the heparin through the bulk material into the fluid. Both effects are, with others, responsible for the burst release, although this phenomenon is still not entirely understood in its full complexity [37]. After an initial large bolus of released drug, the release rate reaches a more stable profile. It could be shown that the amount of eluted drug from the unpolymerized samples increased with higher drug loads; for ewxample, after 144 h of dissolution, a drug release of 30–70% was found. This was also valid for the polymerized matrices. According to this, the release rate of heparin from these different hybrid matrixes increased with higher drug loads of between 2.7 and 8.0 wt.% in gels dried to constant weight and decreased with higher polymerization degrees of between 0% and 71%. As the polymerization degree has a strong influence on the release rate, it could be demonstrated that this organic network causes strong structural changes in the matrix, e.g. cross-linkage, density and hence porosity (Fig. 5). The release kinetics of drugs dissolved from the sol–gelderived matrices can be described by the Higuchi-equation [38,39], which can be taken to show that the released amount of drug per surface area is a function of time (Eq. (6)): rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi pffi De Q¼  ð2 A  eC s Þ  C s  t ¼ k H  t ð6Þ s where Q is the cumulative amount of released drug per surface area (mol m2), D is the diffusion coefficient of the drug in the release medium (m2 s1), e is the porosity, s is the bending of pores, A is the drug load (mol m3), Cs is the saturation dissolubility (mol m3), t is time (s) and kH is the composition-dependent proportionality constant. The Higuchi equation can be used to describe the release kinetics of heparin from P-MA–PS matrices, so the experimental data of eluted drug can be plotted against the square root of time (Fig. 6a–c), followed by linear regression. The errors of the respective gradients (Table 1) can be taken for an estimation of the suitability of the Higuchi equation for P-MA–PS matrices. The errors of the gradients (Table 1) decrease with smaller amounts of cumulative released drug and increase with larger quantities of eluted heparin. However, relative errors are all in the range 2–7%, which shows, that the Higuchi model can be used quite well to describe the release kinetics of heparin from differently drug-loaded and polymerized P-MA–PS matrices. These results are in accordance with other studies on sol–gel derived silica xerogels as drugrelease systems for heparin. The release profiles described in the literature for silica were also fitted to the Higuchi model [24]. The errors (Table 1) can be considered as the degree of the differences between the ideal Higuchi model and the experimental acquired values caused by requirements and simplifications of the model that did not exist during the

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Fig. 6. Plot according to the Higuchi model. The cumulative amount of released heparin from differently drug-loaded samples is plotted against the square root of time followed by linear regression (dashed lines) for matrices with different polymerization degrees; (a) 0%, (b) 36% and (c) 71% polymerization.

dissolution tests. Therefore, the release medium should not underlie swelling or degradation, although both effects occurred marginally during the dissolution experiments from the xerogels. Retrospectively, the relatively good application of the Higuchi model to P-MA–PS drug depots was presumably a result of the very homogeneous porosity of the matrix, which was derived by the sol–gel process and which was also a requirement of the model. 3.5. Release of MAS from the matrix The release of heparin dispersed in an inorganic matrix occurs according to a combined process of diffusion of solvent through the capillarity channels and erosion of the matrix into non-toxic ionic products. The latter phenomenon (biodegradation) takes place by hydrolysis of siloxane bonds into RmSi(OH)4m, which diffuses from the implant into the blood and lymph system and is excreted through the kidneys [31]. For this purpose, it is important to determine the amount of matrix dissolved

in the medium to obtain a first impression of the matrix degradation in vitro. Release of the non-polymerized MAS from the xerogel matrix during the experimental dissolution test was evaluated by measuring dissolved RSi(OH)3 directly by means of UV spectrophotometry due to the C@C signal of the methacrylic functionality in R (=–(CH2)3O(C@O) (C@C)CH3) at 220 nm. In the first 24 h dissolution period unhydrolyzed and unpolymerized molecules near the matrix surface are leached out faster until the dissolution reaches a stable hydrolysis profile. At this point the matrix dissolution becomes linear according to a hydrolysis kinetic of pseudo-zero order because of the large excess of MAS in the system (Fig. 7). The results show that additional organic cross-linkage lowers the degradation of the anorganic matrix substantially, making this effect interesting for controlling the hydrolysis rate. Since all samples underlie the same hydrolysis condition, it can be concluded that the amount of heparin embedded did not have an influence on the degradation

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ferently polymerized and drug-loaded samples was determined as described above and is shown in Table 2. Eluted heparin from the differently polymerized and drug-loaded samples was found to retain its ability to bind to and inactivate thrombin. The results demonstrate that the released heparin retained at least 83% or more of its biological activity. This was valid for each type of tested xerogel and there was no evidence for a significant change in the potencies of the eluted heparin caused by different drug loads or degrees of polymerization of the drug depots. In future works, a further control of the release kinetics of heparin might be achieved by modification of the xerogel matrix with molecules that can bind the drug covalently or ionically and thereby prolong and control the release rate, e.g. by using spacer molecules like aminopropyltrimethoxysilane (APMS) for covalent binding [11] or positively charged polyelectrolytes like polyethylenimine (PEI) for ionic modification [41]. Due to the good adhesion of P-MA–PS matrices onto titanium surfaces and the possibility to adjust material properties as well as the release behaviour of the model drug heparin by many parameters, P-MA–PS might be used as coating for drug-eluting coronary titanium stents.

Fig. 7. Degradation of the organic and inorganic hybrid matrices correlated to different polymerization degrees of the organic network. Each data point is the mean of three samples and error bars are standard deviations.

rate of the matrix, which means that the different drug load does not change the structure of the P-MA–PS matrix. These findings are in accordance with previous results on drug-loaded silica xerogels [40].

4. Conclusion This work shows that organically modified silica ceramics are suitable hybrid matrices for the release of heparin over a prolonged time period. The results demonstrate that drug release from and degradation of the matrix can be controlled by the polymerization degree over a broad spectrum while maintaining the biological activity of the drug. Higher amounts of drug embedded in the matrix showed a faster release profile, which could be described by the Higuchi model. The degradation of the matrix was found to be linear with immersion time, and the amount of embedded heparin did not have an influence on the degradation rate. In conclusion, we have synthesised a novel drug delivery system based on an organically modified silica xerogel system, whereas the release properties can be modified and adjusted according to therapeutic needs.

3.6. Biological activity of heparin Heparin, which inhibits thrombin (T) formation, generates heparin–antithrombin III (ATIII) complex with ATIII in plasma. The active site of heparin responsible for binding to ATIII is the pentasaccaride segment. The biological activity of heparin can be directly measured if an excess amount of thrombin is used to form heparin–ATIII–T complexes. The amount of residual thrombin is determined with Chromozym TH (b-Ala-Gly-Arg-p-nitroanilide diacetate; Sigma). Thrombin acts as a catalyst in splitting of para-nitroaniline (pNA) from Chromozym TH. The pNA release rate was determined by measuring the absorbance at 405 nm. The potency of heparin diffused out of the dif-

Table 2 Biological potencies of heparin eluted from the differently polymerized and drug-loaded samples (poly = polymerization; max = maximal; exp = experimental) Heparin load (mg) 50 Degreepoly (%) Heparin (lg ml1) Activitymax (U ml1) Absorptionexp Activityexp (U ml1) Biological potency (%)

0 0.46 0.080 1.55 0.074 93

100 36 0.41 0.071 1.59 0.072 102

71 0.33 0.058 1.68 0.055 95

0 0.67 0.116 1.24 0.115 99

150 36 0.58 0.100 1.48 0.083 83

71 0.50 0.087 1.48 0.083 95

0 0.99 0.172 0.94 0.153 89

36 0.83 0.144 1.09 0.135 94

71 0.70 0.120 1.15 0.126 105

Aliquots from samples with a drug load of 50 mg were measured 72 h after starting the release experiment; those from the samples with drug loads of 100 and 150 mg were determined in an analogous manner 76 h after the beginning of the drug dissolution.

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