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Precipitation casting of polycaprolactone for applications in tissue engineering and drug delivery
ARTICLE IN PRESS
Biomaterials 25 (2004) 315–325
Precipitation casting of polycaprolactone for applications in tissue engineering and drug delivery A.G.A. Coombesa,*,1, S.C. Rizzib, M. Williamsonc, J.E. Barraletd, S. Downesa,b,2, W.A. Wallacee a
Faculty of Medicine and Health Sciences, Medical School, Queen’s Medical Centre, Nottingham NG7 2UH, UK b Institut fur Biomedizinische Technik, ETH Zurich, University of Zurich, Switzerland c Aston Pharmacy School, Aston University, Aston Triangle, Birmingham B4 7ET, UK d Biomaterials Unit, School of Dentistry, The University of Birmingham, St Chad’s Queensway, Birmingham B4 6NN, UK e Division of Orthopaedics and Accident Surgery, Medical School, Queen’s Medical Centre, Nottingham NG7 2UH, UK Received 17 December 2002; accepted 15 June 2003
Abstract Microporous materials have been produced by gradual precipitation from solutions of poly(e-caprolactone) (PCL) in acetone induced by solvent extraction across a semi-permeable PCL membrane which is formed in situ at the polymer solution/non-solvent interface. Microparticulates of hydroxyapatite and inulin polysaccharide, respectively, were incorporated in precipitation cast PCL matrices to illustrate potential applications in hard tissue repair and macromolecular drug release. Microporous PCL and HA filled PCL materials were found to provide a favourable surface for attachment and growth of primary human osteoblasts in cell culture. The in vitro degradation characteristics of microporous PCL and inulin/PCL materials in PBS at 37 C were monitored over 45 months. Microporous PCL demonstrated zero weight loss, minor changes in molecular weight characteristics and a fairly constant indentation resistance of around 1 MN/m2. Inulin-loaded PCL materials exhibited a total weight loss of approximately 17% after 12 months in PBS. The indentation resistance decreased by 50% from an initial value of 28 MN/m2 in the first 2 months and then remained stable. Precipitation cast materials based on PCL are expected to be useful for formulating long-term, controlled release devices for bioactive molecules such as growth factors and hormones and extended-residence supports for cell growth and tissue development. r 2003 Elsevier Ltd. All rights reserved. Keywords: Poly(e-caprolactone); In vitro degradation; Bone substitute; Drug delivery
1. Introduction Poly(e-caprolactone) (PCL) is a synthetic a-polyester exhibiting a low Tg of around 60 C which imparts a rubbery characteristic to the material. PCL like other members of this family of polymers such as poly(l.lactide) (PLA) and poly(lactide co-glycolide) (PLG), undergoes auto-catalysed bulk hydrolysis. However, the semi-crystalline nature of the polymer extends its *Corresponding author. Fax: +61-29351-4391. E-mail address: [email protected] (A.G.A. Coombes). 1 Present address: Faculty of pharmacy, University of Sydney, NSW 2006, Australia. 2 Present address: Smith and Nephew Group Research Centre, York Science Park, Heslington, York YO1 5DF, UK. 0142-9612/$ - see front matter r 2003 Elsevier Ltd. All rights reserved. doi:10.1016/S0142-9612(03)00535-0
resorption time to over 2 years since the close packed macromolecular arrays retard fluid ingress [1]. The rubbery characteristics of PCL results in high permeability which has been exploited for delivery of low molecular weight drugs such as steroids [2] and vaccines [3]. Copolymerisation of lactic acid and e-caprolactone has been investigated to increase degradation rates and improve processability [2,4]. However, the number of applications for these copolymers has often been limited by poor mechanical properties. In the area of bone repair, PCL has been reinforced with phosphate glass fibres for use as intramedullary, fracture fixation pins [5] and applied in tissue engineered constructs for craniofacial repair [6]. The microstructure and architecture of an implanted device along with surface physicochemical characteristics
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are known to exert profound effects on cell attachment, alignment and proliferation. Implant porosity can allow limited tissue ingrowth for stabilisation of permanent implants or provide pathways for tissue regeneration within or over a biological scaffold or matrix. Despite extensive research, the relationship between pore structure, material properties and the type of tissue ingrowth is still unclear. For example, bony ingrowth predominated in porous polymethylmethacrylate (PMMA) implanted in bone when the pore size was around 450 mm. Connective tissue formed when the pore size was below 100 mm and extensive vascular infiltration was only observed with pores around 1000 mm [7]. In bioceramics a pore size in excess of 100 mm has been defined for effective bone ingrowth [8] to accommodate blood vessel formation with closely associated osteocytes. Structures comprising macropores (150–300 mm) highly interconnected by micropores (o50 mm) have been found to be conducive to ingrowth of fibrocartilaginous tissue in polyurethane meniscal implants [9]. The importance of porous, biodegradable materials for guided tissue repair or for support of seeded cells prior to implantation has stimulated much research into fabrication techniques for 3D matrices or shapeable blocks of porous materials based on the poly(a-hydroxy acids) [10]. These include gel casting using PLA or PLG solutions, variation of the proportion of lactide polymer and particulate fillers to provide voids or pores in the material and extraction of low molecular weight additives such as sodium citrate from the solid polymer. Porous PLG tubes for nerve regeneration have been produced by melt extrusion of a PLG/salt mixture (150– 300 mm crystals) followed by salt leaching, while complex porous scaffolds for bone regeneration have been built up sequentially using solid free-form fabrication [11]. The latter approach offers fine control over the size, orientation and material composition of pores and channels. The initial interactions between cells and biomaterials are usually mediated through cell binding to adsorbed glycoproteins such as fibronectin [12,13]. A number of studies have demonstrated that cell attachment is sensitive to the physicochemical properties of the substrate [14] convincingly so in the case of surface chemistry and topography. Cell development is also responsive to changes in substrate characteristics; materials which promote good cell adhesion, for example, do not necessarily induce cell spreading and migration. Similarly, several groups have reported that surfaces which display the best primary attachment characteristics are not necessarily the substrates on which cell proliferation or differentiation is improved [12,15,16]. The controlled presentation and release of bioactive molecules such as integrin recognition species (e.g. fibronectin) and growth factors (e.g. bone morphogenetic protein (BMP)) by implants is therefore a
key strategy for modulating cell adhesion, migration and function [11,17] so as to optimise tissue organisation and regeneration. Bioceramics such as hydroxyapatite (HA) and tricalcium phosphate (TCP) have often been combined with biodegradable polymers to produce bone substitutes because of their structural similarity to the mineral phase of bone and their osteoconductive and bonebinding properties. This approach also offsets the problems of brittleness and the difficulty of shaping hard ceramic materials to fit bone defects [18,19]. Polypeptide growth factors such as BMP have been included in polymeric carriers in order to create an osteoinductive character [20]. In the present study, particulates of hydroxyapatite and a polysaccharide (inulin) were incorporated in novel PCL microporous matrices produced by a precipitation casting technique. The in vitro degradation behaviour and cell interaction are described to demonstrate the potential for producing bone substitutes and controlled delivery of bioactive macromolecules, respectively.
2. Materials and methods Poly (e-caprolactone) (weight average molecular weight (Mw ) 120,000) was obtained from SolvayInterox, Widnes, UK. Non-sintered hydroxyapatite (HA) powder [P120] consisting of sub-micron particles was obtained from Plasma Biotal, Tideswell, UK. The powder comprises agglomerates of needle-shaped crystals approximately 50–100 nm long and 20 nm wide. The mean Ca/P ratio was determined previously to be 1.65 [18]. Inulin powder (extracted from chicory root, 3–6 mm diameter particles) was obtained from Sigma Chemicals. Production of microporous PCL materials in block form consists of three stages: (1) Casting a solution of PCL in a mould. (2) Addition of a layer of non-solvent to cause polymer precipitation. (3) Extraction of the non-solvent from the hardened polymer with water and/or drying the hardened material. A 12.5% w/v solution of PCL in acetone (4 ml) was prepared by warming to approximately 50 C. The solution was poured into a 10 ml polypropylene (PP) syringe body (used as a mould) and allowed to stand at room temperature for 30 min. Methanol (6 ml) was introduced carefully into the mould to form a layer on top of the PCL solution. Rapid precipitation of the polymer was apparent at the interface between the PCL solution and methanol phase forming a film. After 24 h at room temperature, the sample had hardened by precipitation of the PCL polymer due to acetone
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extraction and could be removed from the mould. The material was immersed in water (40 ml) for 3 days to extract the acetone/methanol non-solvent with a change in medium every 24 h. Samples were finally dried in air at room temperature. PCL films for cell culture experiments were produced in the form of 13 mm diameter discs by casting from a 1% w/ v PCL solution in dichloromethane, onto glass coverslips retained in a 50 mm diameter glass petri dish [18]. 2.1. Hydroxyapatite/PCL materials (HA/PCL) HA powder (500 mg) was mixed with 1 ml acetone to produce a slurry. This was dispersed in 4 ml PCL solution in acetone (12.5% w/v) and added to a 10 ml PP mould. Methanol (5 ml) was added forming a layer on top of the suspension and the sample was retained at room temperature for 4 days to allow complete solidification of the PCL matrix. On demould, the material was immersed in water for 3 days with a change of medium every 24 h. The sample was air dried at room temperature giving rise to a hard, white material. 2.2. Inulin/PCL materials Inulin powder (1 g) was mixed with 4 ml, 12.5% w/v PCL solution in acetone and the suspension was poured into a 10 ml PP mould. Subsequent stages were as described above for HA/PCL materials. 2.3. In vitro degradation studies Samples of microporous PCL materials in the form of disks (9 mm diameter 2.5 mm) and inulin-loaded PCL materials (11 3 mm2) were cut from mouldings using a scalpel. Samples were weighed and incubated in 10 ml PBS (containing 0.02% sodium azide as a bacteriostatic agent) in 50 ml PP tubes at 37 C. At various time periods up to 45 months, at least three samples of unloaded and inulin-loaded PCL materials were removed, rinsed in distilled water and dried at room temperature to constant weight. The thermal properties, molecular weight characteristics, weight loss and indentation resistance of the materials were subsequently monitored to provide a measure of in vitro degradation versus incubation time. 2.4. Molecular weight determination Molecular weight determinations were performed on microporous PCL and inulin/PCL samples before and after incubation in PBS at 37 C for various time periods up to 45 months. The analysis was carried out using a Waters GPC/SEC 810 Baseline system with PIgel Mixed E columns (Polymer Laboratories). Polymer samples were dissolved in tetrahydrofuran to produce a 0.02%
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w/v solution. The inulin/PCL solutions were passed through a 2 mm filter to remove inulin before loading onto the column. The elution rate was 1 ml/min. The GPC system was calibrated using polystyrene molecular weight standards (range 330,000–580 Da) in tetrahydrofuran. Analysis of the weight-average (Mw ), numberaverage (Mn ) molecular weights and polydispersity of the samples was subsequently performed by computerbased correlation of the retention time and concentration of eluted species with the molecular weight calibration curve. 2.5. Thermal analysis The thermal characteristics of microporous PCL, inulin-loaded PCL and the starting polymer in the form of pellets were recorded using a Perkin–Elmer Pyris Diamond differential scanning calorimeter (DSC) fitted with a CCA 7 liquid nitrogen cooling system. Duplicate samples were cooled from 20 C to 90 C at a rate of 100 C/min and held at 90 C for 5 min before heating at a rate of 10 C/min to 100 C. Samples were tested prior to and following immersion in PBS at 37 C for various time periods up to 45 months. The glass transition temperature (Tg ) was taken as the midpoint of the curve between pre- and post-transition baselines. Peak melting temperature (Tm ) and heat of fusion data for the polymer crystalline phase were determined using the software facility of the DSC. The latter measurement was subsequently used to determine the sample percentage crystallinity from the reported heat of fusion of 139.5 J/g for fully crystalline PCL [21]. Indium was used as a standard. 2.6. Indentation resistance The indentation resistance of microporous PCL and inulin-loaded PCL materials was investigated using a CNS Farnell QTS Texture Analyser (CNS Farnell, Borehamwood, Herts, UK). Samples were loaded in compression at a rate of 5 mm/min over a distance of 1.0 mm using a flat-faced circular, stainless steel indentor, 2 mm in diameter. The indentation resistance of the material was calculated from the recorded compressive force at 1 mm sample deflection or the maximum force experienced prior to sample failure and the compressed sample area (MN/m2). Samples were tested prior to and following immersion in PBS at 37 C for various time periods up to 45 months. 2.7. Scanning electron microscopy (SEM) The morphology of PCL and inulin/PCL materials before and after incubation in PBS at 37 C was analysed using a JEOL JSM-5300LV SEM. Liquid nitrogen freeze-fractured samples were attached to aluminium
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SEM stubs and sputter coated with gold prior to examination. HA/PCL materials were examined after fracture at room temperature. 2.8. Osteoblast interactions with PCL and HA/PCL materials A cell culture model was applied to elucidate the characteristics of attachment of primary human osteoblasts to microporous PCL and HA/PCL materials and to provide a measure of the biocompatibility of the two materials. 2.9. Human osteoblast (HOB) cell culture Primary human osteoblasts were isolated from explants of trabecular bone taken from femoral heads at the time of surgery for total joint replacement [22]. Cells were cultured in vitro using Dulbecco’s Modified Eagles Medium (DMEM), supplemented with 10% (v/v) foetal calf serum, 0.02 m HEPES buffer, 150I mg/ml ascorbic acid (BDH, Poole, UK), 2 mm l-Glutamine, 1% Penicillin/streptomycin and 1% (v/v) non-essential amino acids (Gibco, BRL, Paisley, UK). The cells were incubated at 37 C, 5% CO2 and the culture medium was changed every 2 days. Upon confluence, cells were trypsinised by removing the medium, washed using 4 ml sterile phosphate buffered saline (PBS) and incubated at 37 C with 0.02% trypsin solution (Sigma, Poole, UK) in PBS with 1% HEPES. After 10 min, the cells were removed from the flasks and resuspended in complete medium. A cell viability count was performed using the trypan blue exclusion stain. 2.10. Osteoblast interactions with PCL and HA/PCL surfaces Disc-shaped specimens of microporous PCL and HA/ PCL materials were washed in 70% and 100% ethanol and dried at 37 C for 2 h under aseptic conditions immediately prior to use in tissue culture. Samples were inserted in 48-well plates and cells were seeded at a concentration of 50,000 cells/well in complete medium at 37 C and 5% CO2 : Incubation times of 90 min, 4 and 24 h, respectively, were investigated. PCL films were tested in 24-well plates using 100,000 cells/well for 90 min and 4 h experiments and 50,000 cells/well for 24 h experiments. Tissue culture plastic (Thermanox) was used as a control surface. 2.11. Investigation of cell morphology using environmental scanning electron microscopy (ESEM) The morphology of HOB cells on the PCL and HA/ PCL materials after 90 min, 4 and 24 h in cell culture was investigated using a Philips ESEM (PEG XL-30)
operated at 5–10 kV in a sample chamber vacuum between 2 and 4 Torr. Following incubation, the culture medium was removed and the samples were rinsed once with 0.1 m PBS. The cells were fixed using 1.5% glutaraldehyde in 0.1 m PBS for 30 min at 4 C. After removing the fixative the cells were rinsed twice with 0.1 m PBS and post-fixed in osmium tetroxide (2% w/v osmium tetroxide in 0.1 m PBS) for 30 min at room temperature. After a further rinse in distilled water the samples were dehydrated through a graded series of ethanol (50%,70%,90%,100%) for 2 5 min. The dehydration was completed in Hexamethyl-disilazane (HMDS). Uncoated samples were used for ESEM analysis of cell/biomaterial interaction.
3. Results Addition of methanol to PCL solution results in rapid polymer precipitation to form an interfacial film which serves both to stabilise the interface and provide a semipermeable membrane for exchange of solvent and nonsolvent. Gradual crystallisation subsequently occurs along a front proceeding from the PCL solution/ methanol interface towards the base of the mould. In the absence of methanol, a viscous fluid was obtained when PCL solution was allowed to stand at room temperature overnight. Use of a 12.5% w/v solution of PCL in a good solvent (dichloromethane) did not result in precipitation. In addition, the replacement of methanol by water did not result in the formation of two distinct phases. Instead, a fine precipitate of PCL collected at the base of the mould. The minimum PCL solution concentration required to produce blocks of material was found to be in the region of 10% w/v. The use of solution concentrations below this value resulted in formation of a fine suspension of precipitates. A critical PCL solution concentration is necessary to produce an effective chain entanglement density which subsequently results in robust, cohesive materials. For example, the dried material prepared using a 10% w/v PCL solution tended to fragment fairly easily on handling. On drying, a white, uniform, cylindrical material was obtained, with an absence of voids and cracking (visual assessment). The density of microporous PCL determined by weighing 2 mm thick discs was 0.2970.002 g/ cm3. Examination of the material by SEM revealed a surface comprising flat, smooth areas and fibrous regions. The underlying morphology was produced by an array of crystalline, lamellar structures giving rise to pores of irregular size and shape of the order of 1–10 mm dimensions (Fig. 1A). No changes in morphology were evident after incubation of samples for 45 months in PBS at 37 C.
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Fig. 1. Scanning electron micrographs: (A) Microporous PCL; (B) HA/PCL material and (C) Inulin/PCL material.
The density of HA/PCL material was 0.4970.003 g/ cm3. SEM examination revealed the irregular, porous morphology formed by the PCL lamellar structures (Fig. 1B). HA particulates less than 5 mm in size were evident, dispersed and embedded within the PCL phase. The density of inulin/PCL material was 0.8970.002 g/ cm3. SEM examination revealed agglomerates of roughly spherical inulin particles, approximately 5 mm in size (arrowed in Fig. 1C) enmeshed in a PCL matrix. No changes in morphology were evident after incubation of samples for 45 months in PBS at 37 C. 3.1. Weight loss characteristics Weight loss was not recorded for microporous PCL samples in PBS at 37 C over the 45 month time course of the study. This is expected on the basis of past work which revealed that significant weight loss of PCL implants does not occur until the molecular weight falls to around 5000 [21]. The inulin/PCL materials exhibited a mean weight loss of 10.2% after 1 month in PBS and 17.6% after 12 months. No further weight loss was recorded up to 45 months.
3.2. Molecular weight changes The molecular weight characteristics of microporous PCL and inulin/PCL materials pre- and post-incubation in PBS at 37 C are presented in Table 1. Relatively minor changes were measured in molecular weight and differences in molecular weight, respectively, between each sample type over 12 months, reflecting the slow degradation rate of PCL. At extended degradation times of 32 and 45 months, microporous PCL tends to exhibit a more rapid decrease in molecular weight than the inulin/PCL possibly due to a higher porosity of the unloaded material (Fig. 1) which both facilitates permeation of the PBS incubation medium and increases the material surface area for hydrolysis. 3.3. Thermal analysis The melting point and Tg of microporous PCL and inulin/PCL materials were, in general, close to the values usually quoted for PCL i.e. 60 C and 60 C, respectively (Table 2). The higher value of Tg recorded for microporous PCL incubated for 45 months in PBS
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Table 1 The molecular weight characteristics of microporous PCL and inulin/ PCL materials pre- and post-incubation in PBS at 37 C Sample
Incubation time (months)
Mn
Mw
Polydispersity
PCL
0 1 2 3 6 12 32 45
78,778 77,999 77,406 75,040 75,550 74,112 55,679 40,137
113,508 112,887 112,090 109,153 110,442 108,011 87,255 79,710
1.44 1.45 1.45 1.46 1.46 1.46 1.57 1.99
0
76,234
110,610
1.45
1 2 4 9 12 32 45
71,713 72,653 75,291 69,164 68,274 60,397 57,182
105,985 105,349 108,878 105,616 102,708 99,933 94,831
1.48 1.45 1.45 1.50 1.50 1.66 1.66
Inulin/ PCL
Table 2 The thermal characteristics of microporous PCL and inulin/PCL materials pre- and post-incubation in PBS at 37 C Tg ( C)
Sample
Incubation time (months)
Tm ( C)
PCL
0 6 12 32 45
60.3 60.6 62.5 61.4 61.8
61.6 63.5 62.0 59.2 55.9
77.9 78.0 78.6 77.4 79.8
Inulin/PCL
0 4 12 30 44
61.8 62.7 62.2 62.6 61.9
59.0 61.7 58.8 58.3 57.9
61.6 62.0 55.9 56.2 54.2
55.3
62.8
59.0
Unprocessed PCL pellets
Crystallinity (%)
suggests a constraint on chain mobility in the amorphous phase. High crystallinity of approximately 80% developed in microporous PCL during production and remained unchanged during incubation in PBS at 37 C. In contrast the crystallinity of inulin/PCL materials was around 25% lower (similar to unprocessed pellets) and also remained unchanged during incubation in PBS. 3.4. Indentation resistance Microporous PCL materials are resilient under small strains but can be compressed readily resulting in permanent deformation. Values of the indentation
Table 3 The indentation resistance of microporous PCL and inulin/PCL materials following incubation in PBS at 37 C Immersion time in PBS (months)
0 1 2 3 4 6 9 12 32 45
Indentation resistance (mn/m2) PCL
Inulin/PCL
1.270.3 1.270.2 0.970.5 1.170.5 ND 1.170.1 ND 1.170.03 1.070.1 1.270.1
24.076.0 22.172.1 10.970.5 ND 10.470.8 ND 11.270.7 11.673.6 9.674.0 10.070.2
resistance of microporous PCL and inulin/PCL materials, pre- and post-immersion in PBS at 37 C are listed in Table 3. The indentation resistance of microporous PCL was virtually unchanged after 45 months immersion in PBS at 37 C at around 1.2 MN/m2. The force/deflection curve also remained unchanged by immersion of samples in PBS, consisting of a gradual, almost linear increase in load with increasing material compaction and density below the indentor probe (Fig. 2D). The minor changes recorded in indentation behaviour correlate with the hydrolytic stability of PCL as indicated by molecular weight determinations (Table 1). Although differences in test methods preclude strict comparisons, the indentation resistance of microporous PCL is similar to the compressive strength (1.0 MN/m2) of microporous PLA/PLG blends prepared by gel casting [23]. The inulin/PCL materials are hard and rigid materials due to the high (67% w/w) loading of inulin powder. This gives rise to an indentation resistance which is almost 20 times higher than microporous PCL (Table 3). In contrast to microporous PCL a major decrease in compression resistance of almost 70% was measured for inulin/PCL materials after 2 months immersion in PBS at 37 C which then remained fairly constant over the next 43 months. The reduction in mechanical properties of the inulin/PCL materials coincides with a loss in sample weight of 11% due to dissolution and extraction of the reinforcing polysaccharide phase. Distinct changes in the indentation force/deflection curve were observed with time of incubation in PBS. Up to 12 months, a two-stage curve was apparent exhibiting a yield point followed by an almost linear load/deflection trace of lower slope (Fig. 2A). This behaviour is similar to that reported for microporous gel-cast materials produced from PLA, PLG and their blends [23]. After 12 months, inulin/PCL materials gave rise to a yield point but fractured before the test limit of deflection
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resulting in radial crack propagation and a gradual decrease in load-bearing ability (Fig. 2B). An increase in incubation time to 32 months and above resulted in fracture of inulin/PCL samples without yielding, gradual crack propagation and decreased load-bearing ability (Fig. 2C). 3.5. Osteoblast-interaction with PCL and PCL/HA materials
extraction may be applicable such as PLA or high llactide-containing PLG copolymers. Various materials may be added to the PCL solution to adjust the properties of the finished material such as degradation rate, morphology and the pattern of drug release, or to modulate cell interaction with the material. The absence of any weight loss for microporous PCL materials is in line with the findings of Pitt et al. [21] who investigated the degradation of implanted, sub-dermal, drug delivery devices produced from PCL (Mn 50,000) over a timescale of 52 months. Weight loss was not recorded until after approximately 2.5 years when the molecular weight (Mn ) had fallen to 5000, reflecting the requirement for degraded chain fragments to be below a critical chain length for diffusion from the matrix. PCL is a semi-crystalline aliphatic polyester. As such the degree of crystallinity is important in controlling degradation since the crystalline regions retard fluid ingress. The method of manufacture of microporous
60 A
40 B
Force (N)
The characteristic HOB cell morphology observed on the various substrates after 90 min, 4 and 24 h in cell culture is indicated in Table 4 and illustrated in Figs. 3 and 4. The degree of cell spreading was higher on Thermanox disks than on PCL films at time points of 90 min and 4 h and equivalent at 24 h. The notable differences in cell morphology on microporous PCL and HA/PCL materials compared with PCL films and Thermanox relates to the coalescence of attached cells after 4 h in culture. Individual cells were still discernible on Thermanox up to 24 h (Fig. 3). In contrast, Fig. 4D shows HOB cell layers draped over the surface features of HA/PCL material after 4 h, masking the underlying grainy texture. The tops of rounded cells assist in identifying the areas of cell coalescence. In Fig. 4B a sheet of HOB cells formed after 24 h in culture spans the underlying fibrous surface structure of microporous PCL.
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4. Discussion Blocks of microporous materials have been formed by gradual precipitation from solutions of PCL in acetone induced by solvent extraction across a semi-permeable membrane which is formed in situ at a PCL solution/ methanol interface. The viscosity of PCL solutions above 10% w/v concentration is sufficiently high to resist mixing with the non-solvent phase. This behaviour coupled with rapid production of an interfacial film results in the separation of two distinct fluid layers. Solvent exchange across the film subsequently induces gradual precipitation of PCL to form a block of microporous material. Although PCL provided the focus for the present work, other polymers which show a tendency to crystallise from solution on solvent
C 20
D 0.5
0
1
Deflection (mm) Fig. 2. Characteristic indentation force–deflection curves for inulin/ PCL materials (A,B,C) and microporous PCL following incubation in PBS at 37 C: (A) inulin/PCL, 0–12 months in PBS; (B) inulin/PCL, 12–32 months in PBS; (C) inulin/PCL, >32 months in PBS and (D) microporous PCL.
Table 4 The morphology of primary human osteoblasts cultured on tissue culture plastic (Thermanox), PCL films and microporous PCL materials Time point
90 min 4h 24 h
Appearance of HOB cells Microporous PCL and HA/PCL
PCL film
Thermanox
Rounded, cell processes visible Spread, forming cell layer Cell layer, individual cells not distinguishable
Rounded Rounded and spread cells Spread cells
Rounded and spread cells Spread cells Individual cells discernible
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(A)
(B)
RC
SC
(C)
SC
Fig. 3. The morphology of human osteoblasts following cell culture on PCL films and Thermanox. (A) Rounded cells (RC) on PCL film after 90 min in cell culture; (B) Rounded and spread cells (SC) on Thermanox after 90 min and (C) Spread cells on Thermanox after 4 h in cell culture.
PCL involving slow, gradual precipitation from solution allows the development of high levels of crystallinity similar to the behaviour of solvent-cast PLA films [23]. This essentially restricts the scope for further crystallization during incubation in PBS (Table 2). In contrast Pitt et al. [21] measured a rapid increase in crystallinity of PCL drug delivery devices over the first 4 weeks postimplantation from 45% to 50% which was ascribed to annealing of the polymer at body temperature. The crystallinity gradually increased to a value of approximately 80% after 30 months which was attributed to crystallization of tie chain segments from the amorphous phase following chain cleavage, facilitated by the low Tg of PCL. The relative insensitivity of the crystallinity of inulin/PCL materials to incubation in PBS in the present study (Table 2) suggests that the PCL chain segments in the amorphous phase are strongly bound to the polysaccharide phase thereby restricting the molecular mobility necessary for reorganization and crystallization. The constancy of mechanical performance of microporous PCL samples following immersion in PBS at 37 C for 45 months suggests that the reduction in fracture toughness observed for inulin/PCL materials at
and beyond 12 months is due to changes in the structural stability of inulin. This behaviour may be exploited to facilitate the breakdown of inulin-loaded implants in vivo by tissue ingrowth, thereby enhancing the rate of elimination of implanted material by phagocytosis. Osteoblasts display distinct morphological changes on biomaterials, often within 2 h of cell culture in the case of HA and alumina surfaces [24]. Cytoskeletal reorganisation (cell flattening/spreading and process extension) is anticipated by 4 h following contact of rounded cells with the surface while cells are expected to be flat, spread and to have started proliferating after 24 h [25]. The timing of the above morphological changes is not exact and varies with substrate characteristics and cell type [16,25,26]. Spreading of HOB cells occurred on microporous PCL and HA/PCL materials in the same time period as Thermanox which is surface modified to promote favourable cell attachment and growth characteristics. HOB cells were found to be spread to a larger extent on microporous PCL and HA/PCL materials at 4 h than on PCL films. These differences may be partly explained in terms of competitive and/or selective adsorption of cell adhesion molecules, such as fibronec-
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(B)
(A)
P
323
P CL D CL
(D)
(C)
CL CL
CL
RC
Fig. 4. The morphology of human osteoblasts following cell culture on microporous PCL and HA/PCL materials. (A) Microporous PCL after 90 min in cell culture. (B) Microporous PCL after 24 h. CL cell layer, P-polymer, D-identifying dehydration cracks. (C) HA/PCL material after 90 min in cell culture. (D) HA/PCL after 4 h.
tin, to produce a favourable packing density and conformation-controlled presentation of adhesion sequences. HA is known to enhance cell adhesion and spreading probably through an inherently high capacity for adsorbing proteins [27]. In addition, the surface topography of the microporous PCL materials may be more conducive to cell attachment and development than a smooth film surface. The numbers of adhered cells, cell shape and attachment forces have been shown by many investigators [13,18] to be affected by surface morphology. The propensity for cell coalescence to form sheets on the microporous PCL and HA/PCL materials (Fig. 4) rather than separated cells indicates the presence of a high density of anchorage sites for integrin binding leading to extensive spreading and cell–cell contacts. This in turn may positively influence the pattern of mechanical stresses experienced by the cytoskeleton [28]. Cell coalescence on PCL and HA/PCL materials may also favour formation of the osteoblast sheets considered advantageous for correct development of bone tissue [29]. The low resorption rate of precipitation cast materials based on PCL recommends their use as long-term,
delivery systems for bioactive molecules such as growth factors and hormones and extended-residence supports for cell growth and tissue development. Growth factors such as BMP may be incorporated in PCL microporous materials by simple admixing of spray-dried particulates, thereby providing a bone graft substitute which both supports and stimulates local bone growth to achieve integration with host tissue. The potential for controlling drug release from PCL matrices produced by precipitation casting was evaluated using inulin polysaccharide (Mw 5000) as a model biopharmaceutical. Inulin has been used to investigate the release of macromolecules from collagen gels [30] since inulin is filtered by the kidneys and excreted rather than being bound to plasma proteins or metabolized. Inulin release from a drug delivery device in vivo may thus be directly related to the excretion rate. Although the duration of inulin release from PCL matrices was extended to 12 months, the cumulative release was limited to approximately 18% w/w, indicating that the polysaccharide phase is efficiently coated by PCL and pore connectivity is poorly developed. Water soluble polymers are frequently incorporated in polymeric films and
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microspheres to modulate drug delivery [31]. This approach may be useful for improving the rate and duration of macromolecule release from microporous PCL matrices. Copolymerisation and polymer blending provide alternative approaches via control of the resorption characteristics of the matrix.
5. Conclusions Precipitation casting presents a novel approach for producing microporous polymeric, block-form materials and particulate-loaded compositions. The low resorption rate recommends application of PCLbased castings for long-term, release of bioactive molecules and extended-residence cell supports in tissue engineering.
Acknowledgements Thanks are due to Tony Evans, Solvay Interox, UK for providing the PCL molecular weight data in Table 1. AGAC acknowledges the prior financial support of the Division of Orthopaedics and Accident Surgery, Queen’s Medical Centre, University of Nottingham.
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Precipitation casting of polycaprolactone for applications in tissue engineering and drug delivery A.G.A. Coombesa,*,1, S.C. Rizzib, M. Williamsonc, J.E. Barraletd, S. Downesa,b,2, W.A. Wallacee a
Faculty of Medicine and Health Sciences, Medical School, Queen’s Medical Centre, Nottingham NG7 2UH, UK b Institut fur Biomedizinische Technik, ETH Zurich, University of Zurich, Switzerland c Aston Pharmacy School, Aston University, Aston Triangle, Birmingham B4 7ET, UK d Biomaterials Unit, School of Dentistry, The University of Birmingham, St Chad’s Queensway, Birmingham B4 6NN, UK e Division of Orthopaedics and Accident Surgery, Medical School, Queen’s Medical Centre, Nottingham NG7 2UH, UK Received 17 December 2002; accepted 15 June 2003
Abstract Microporous materials have been produced by gradual precipitation from solutions of poly(e-caprolactone) (PCL) in acetone induced by solvent extraction across a semi-permeable PCL membrane which is formed in situ at the polymer solution/non-solvent interface. Microparticulates of hydroxyapatite and inulin polysaccharide, respectively, were incorporated in precipitation cast PCL matrices to illustrate potential applications in hard tissue repair and macromolecular drug release. Microporous PCL and HA filled PCL materials were found to provide a favourable surface for attachment and growth of primary human osteoblasts in cell culture. The in vitro degradation characteristics of microporous PCL and inulin/PCL materials in PBS at 37 C were monitored over 45 months. Microporous PCL demonstrated zero weight loss, minor changes in molecular weight characteristics and a fairly constant indentation resistance of around 1 MN/m2. Inulin-loaded PCL materials exhibited a total weight loss of approximately 17% after 12 months in PBS. The indentation resistance decreased by 50% from an initial value of 28 MN/m2 in the first 2 months and then remained stable. Precipitation cast materials based on PCL are expected to be useful for formulating long-term, controlled release devices for bioactive molecules such as growth factors and hormones and extended-residence supports for cell growth and tissue development. r 2003 Elsevier Ltd. All rights reserved. Keywords: Poly(e-caprolactone); In vitro degradation; Bone substitute; Drug delivery
1. Introduction Poly(e-caprolactone) (PCL) is a synthetic a-polyester exhibiting a low Tg of around 60 C which imparts a rubbery characteristic to the material. PCL like other members of this family of polymers such as poly(l.lactide) (PLA) and poly(lactide co-glycolide) (PLG), undergoes auto-catalysed bulk hydrolysis. However, the semi-crystalline nature of the polymer extends its *Corresponding author. Fax: +61-29351-4391. E-mail address: [email protected] (A.G.A. Coombes). 1 Present address: Faculty of pharmacy, University of Sydney, NSW 2006, Australia. 2 Present address: Smith and Nephew Group Research Centre, York Science Park, Heslington, York YO1 5DF, UK. 0142-9612/$ - see front matter r 2003 Elsevier Ltd. All rights reserved. doi:10.1016/S0142-9612(03)00535-0
resorption time to over 2 years since the close packed macromolecular arrays retard fluid ingress [1]. The rubbery characteristics of PCL results in high permeability which has been exploited for delivery of low molecular weight drugs such as steroids [2] and vaccines [3]. Copolymerisation of lactic acid and e-caprolactone has been investigated to increase degradation rates and improve processability [2,4]. However, the number of applications for these copolymers has often been limited by poor mechanical properties. In the area of bone repair, PCL has been reinforced with phosphate glass fibres for use as intramedullary, fracture fixation pins [5] and applied in tissue engineered constructs for craniofacial repair [6]. The microstructure and architecture of an implanted device along with surface physicochemical characteristics
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are known to exert profound effects on cell attachment, alignment and proliferation. Implant porosity can allow limited tissue ingrowth for stabilisation of permanent implants or provide pathways for tissue regeneration within or over a biological scaffold or matrix. Despite extensive research, the relationship between pore structure, material properties and the type of tissue ingrowth is still unclear. For example, bony ingrowth predominated in porous polymethylmethacrylate (PMMA) implanted in bone when the pore size was around 450 mm. Connective tissue formed when the pore size was below 100 mm and extensive vascular infiltration was only observed with pores around 1000 mm [7]. In bioceramics a pore size in excess of 100 mm has been defined for effective bone ingrowth [8] to accommodate blood vessel formation with closely associated osteocytes. Structures comprising macropores (150–300 mm) highly interconnected by micropores (o50 mm) have been found to be conducive to ingrowth of fibrocartilaginous tissue in polyurethane meniscal implants [9]. The importance of porous, biodegradable materials for guided tissue repair or for support of seeded cells prior to implantation has stimulated much research into fabrication techniques for 3D matrices or shapeable blocks of porous materials based on the poly(a-hydroxy acids) [10]. These include gel casting using PLA or PLG solutions, variation of the proportion of lactide polymer and particulate fillers to provide voids or pores in the material and extraction of low molecular weight additives such as sodium citrate from the solid polymer. Porous PLG tubes for nerve regeneration have been produced by melt extrusion of a PLG/salt mixture (150– 300 mm crystals) followed by salt leaching, while complex porous scaffolds for bone regeneration have been built up sequentially using solid free-form fabrication [11]. The latter approach offers fine control over the size, orientation and material composition of pores and channels. The initial interactions between cells and biomaterials are usually mediated through cell binding to adsorbed glycoproteins such as fibronectin [12,13]. A number of studies have demonstrated that cell attachment is sensitive to the physicochemical properties of the substrate [14] convincingly so in the case of surface chemistry and topography. Cell development is also responsive to changes in substrate characteristics; materials which promote good cell adhesion, for example, do not necessarily induce cell spreading and migration. Similarly, several groups have reported that surfaces which display the best primary attachment characteristics are not necessarily the substrates on which cell proliferation or differentiation is improved [12,15,16]. The controlled presentation and release of bioactive molecules such as integrin recognition species (e.g. fibronectin) and growth factors (e.g. bone morphogenetic protein (BMP)) by implants is therefore a
key strategy for modulating cell adhesion, migration and function [11,17] so as to optimise tissue organisation and regeneration. Bioceramics such as hydroxyapatite (HA) and tricalcium phosphate (TCP) have often been combined with biodegradable polymers to produce bone substitutes because of their structural similarity to the mineral phase of bone and their osteoconductive and bonebinding properties. This approach also offsets the problems of brittleness and the difficulty of shaping hard ceramic materials to fit bone defects [18,19]. Polypeptide growth factors such as BMP have been included in polymeric carriers in order to create an osteoinductive character [20]. In the present study, particulates of hydroxyapatite and a polysaccharide (inulin) were incorporated in novel PCL microporous matrices produced by a precipitation casting technique. The in vitro degradation behaviour and cell interaction are described to demonstrate the potential for producing bone substitutes and controlled delivery of bioactive macromolecules, respectively.
2. Materials and methods Poly (e-caprolactone) (weight average molecular weight (Mw ) 120,000) was obtained from SolvayInterox, Widnes, UK. Non-sintered hydroxyapatite (HA) powder [P120] consisting of sub-micron particles was obtained from Plasma Biotal, Tideswell, UK. The powder comprises agglomerates of needle-shaped crystals approximately 50–100 nm long and 20 nm wide. The mean Ca/P ratio was determined previously to be 1.65 [18]. Inulin powder (extracted from chicory root, 3–6 mm diameter particles) was obtained from Sigma Chemicals. Production of microporous PCL materials in block form consists of three stages: (1) Casting a solution of PCL in a mould. (2) Addition of a layer of non-solvent to cause polymer precipitation. (3) Extraction of the non-solvent from the hardened polymer with water and/or drying the hardened material. A 12.5% w/v solution of PCL in acetone (4 ml) was prepared by warming to approximately 50 C. The solution was poured into a 10 ml polypropylene (PP) syringe body (used as a mould) and allowed to stand at room temperature for 30 min. Methanol (6 ml) was introduced carefully into the mould to form a layer on top of the PCL solution. Rapid precipitation of the polymer was apparent at the interface between the PCL solution and methanol phase forming a film. After 24 h at room temperature, the sample had hardened by precipitation of the PCL polymer due to acetone
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extraction and could be removed from the mould. The material was immersed in water (40 ml) for 3 days to extract the acetone/methanol non-solvent with a change in medium every 24 h. Samples were finally dried in air at room temperature. PCL films for cell culture experiments were produced in the form of 13 mm diameter discs by casting from a 1% w/ v PCL solution in dichloromethane, onto glass coverslips retained in a 50 mm diameter glass petri dish [18]. 2.1. Hydroxyapatite/PCL materials (HA/PCL) HA powder (500 mg) was mixed with 1 ml acetone to produce a slurry. This was dispersed in 4 ml PCL solution in acetone (12.5% w/v) and added to a 10 ml PP mould. Methanol (5 ml) was added forming a layer on top of the suspension and the sample was retained at room temperature for 4 days to allow complete solidification of the PCL matrix. On demould, the material was immersed in water for 3 days with a change of medium every 24 h. The sample was air dried at room temperature giving rise to a hard, white material. 2.2. Inulin/PCL materials Inulin powder (1 g) was mixed with 4 ml, 12.5% w/v PCL solution in acetone and the suspension was poured into a 10 ml PP mould. Subsequent stages were as described above for HA/PCL materials. 2.3. In vitro degradation studies Samples of microporous PCL materials in the form of disks (9 mm diameter 2.5 mm) and inulin-loaded PCL materials (11 3 mm2) were cut from mouldings using a scalpel. Samples were weighed and incubated in 10 ml PBS (containing 0.02% sodium azide as a bacteriostatic agent) in 50 ml PP tubes at 37 C. At various time periods up to 45 months, at least three samples of unloaded and inulin-loaded PCL materials were removed, rinsed in distilled water and dried at room temperature to constant weight. The thermal properties, molecular weight characteristics, weight loss and indentation resistance of the materials were subsequently monitored to provide a measure of in vitro degradation versus incubation time. 2.4. Molecular weight determination Molecular weight determinations were performed on microporous PCL and inulin/PCL samples before and after incubation in PBS at 37 C for various time periods up to 45 months. The analysis was carried out using a Waters GPC/SEC 810 Baseline system with PIgel Mixed E columns (Polymer Laboratories). Polymer samples were dissolved in tetrahydrofuran to produce a 0.02%
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w/v solution. The inulin/PCL solutions were passed through a 2 mm filter to remove inulin before loading onto the column. The elution rate was 1 ml/min. The GPC system was calibrated using polystyrene molecular weight standards (range 330,000–580 Da) in tetrahydrofuran. Analysis of the weight-average (Mw ), numberaverage (Mn ) molecular weights and polydispersity of the samples was subsequently performed by computerbased correlation of the retention time and concentration of eluted species with the molecular weight calibration curve. 2.5. Thermal analysis The thermal characteristics of microporous PCL, inulin-loaded PCL and the starting polymer in the form of pellets were recorded using a Perkin–Elmer Pyris Diamond differential scanning calorimeter (DSC) fitted with a CCA 7 liquid nitrogen cooling system. Duplicate samples were cooled from 20 C to 90 C at a rate of 100 C/min and held at 90 C for 5 min before heating at a rate of 10 C/min to 100 C. Samples were tested prior to and following immersion in PBS at 37 C for various time periods up to 45 months. The glass transition temperature (Tg ) was taken as the midpoint of the curve between pre- and post-transition baselines. Peak melting temperature (Tm ) and heat of fusion data for the polymer crystalline phase were determined using the software facility of the DSC. The latter measurement was subsequently used to determine the sample percentage crystallinity from the reported heat of fusion of 139.5 J/g for fully crystalline PCL [21]. Indium was used as a standard. 2.6. Indentation resistance The indentation resistance of microporous PCL and inulin-loaded PCL materials was investigated using a CNS Farnell QTS Texture Analyser (CNS Farnell, Borehamwood, Herts, UK). Samples were loaded in compression at a rate of 5 mm/min over a distance of 1.0 mm using a flat-faced circular, stainless steel indentor, 2 mm in diameter. The indentation resistance of the material was calculated from the recorded compressive force at 1 mm sample deflection or the maximum force experienced prior to sample failure and the compressed sample area (MN/m2). Samples were tested prior to and following immersion in PBS at 37 C for various time periods up to 45 months. 2.7. Scanning electron microscopy (SEM) The morphology of PCL and inulin/PCL materials before and after incubation in PBS at 37 C was analysed using a JEOL JSM-5300LV SEM. Liquid nitrogen freeze-fractured samples were attached to aluminium
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SEM stubs and sputter coated with gold prior to examination. HA/PCL materials were examined after fracture at room temperature. 2.8. Osteoblast interactions with PCL and HA/PCL materials A cell culture model was applied to elucidate the characteristics of attachment of primary human osteoblasts to microporous PCL and HA/PCL materials and to provide a measure of the biocompatibility of the two materials. 2.9. Human osteoblast (HOB) cell culture Primary human osteoblasts were isolated from explants of trabecular bone taken from femoral heads at the time of surgery for total joint replacement [22]. Cells were cultured in vitro using Dulbecco’s Modified Eagles Medium (DMEM), supplemented with 10% (v/v) foetal calf serum, 0.02 m HEPES buffer, 150I mg/ml ascorbic acid (BDH, Poole, UK), 2 mm l-Glutamine, 1% Penicillin/streptomycin and 1% (v/v) non-essential amino acids (Gibco, BRL, Paisley, UK). The cells were incubated at 37 C, 5% CO2 and the culture medium was changed every 2 days. Upon confluence, cells were trypsinised by removing the medium, washed using 4 ml sterile phosphate buffered saline (PBS) and incubated at 37 C with 0.02% trypsin solution (Sigma, Poole, UK) in PBS with 1% HEPES. After 10 min, the cells were removed from the flasks and resuspended in complete medium. A cell viability count was performed using the trypan blue exclusion stain. 2.10. Osteoblast interactions with PCL and HA/PCL surfaces Disc-shaped specimens of microporous PCL and HA/ PCL materials were washed in 70% and 100% ethanol and dried at 37 C for 2 h under aseptic conditions immediately prior to use in tissue culture. Samples were inserted in 48-well plates and cells were seeded at a concentration of 50,000 cells/well in complete medium at 37 C and 5% CO2 : Incubation times of 90 min, 4 and 24 h, respectively, were investigated. PCL films were tested in 24-well plates using 100,000 cells/well for 90 min and 4 h experiments and 50,000 cells/well for 24 h experiments. Tissue culture plastic (Thermanox) was used as a control surface. 2.11. Investigation of cell morphology using environmental scanning electron microscopy (ESEM) The morphology of HOB cells on the PCL and HA/ PCL materials after 90 min, 4 and 24 h in cell culture was investigated using a Philips ESEM (PEG XL-30)
operated at 5–10 kV in a sample chamber vacuum between 2 and 4 Torr. Following incubation, the culture medium was removed and the samples were rinsed once with 0.1 m PBS. The cells were fixed using 1.5% glutaraldehyde in 0.1 m PBS for 30 min at 4 C. After removing the fixative the cells were rinsed twice with 0.1 m PBS and post-fixed in osmium tetroxide (2% w/v osmium tetroxide in 0.1 m PBS) for 30 min at room temperature. After a further rinse in distilled water the samples were dehydrated through a graded series of ethanol (50%,70%,90%,100%) for 2 5 min. The dehydration was completed in Hexamethyl-disilazane (HMDS). Uncoated samples were used for ESEM analysis of cell/biomaterial interaction.
3. Results Addition of methanol to PCL solution results in rapid polymer precipitation to form an interfacial film which serves both to stabilise the interface and provide a semipermeable membrane for exchange of solvent and nonsolvent. Gradual crystallisation subsequently occurs along a front proceeding from the PCL solution/ methanol interface towards the base of the mould. In the absence of methanol, a viscous fluid was obtained when PCL solution was allowed to stand at room temperature overnight. Use of a 12.5% w/v solution of PCL in a good solvent (dichloromethane) did not result in precipitation. In addition, the replacement of methanol by water did not result in the formation of two distinct phases. Instead, a fine precipitate of PCL collected at the base of the mould. The minimum PCL solution concentration required to produce blocks of material was found to be in the region of 10% w/v. The use of solution concentrations below this value resulted in formation of a fine suspension of precipitates. A critical PCL solution concentration is necessary to produce an effective chain entanglement density which subsequently results in robust, cohesive materials. For example, the dried material prepared using a 10% w/v PCL solution tended to fragment fairly easily on handling. On drying, a white, uniform, cylindrical material was obtained, with an absence of voids and cracking (visual assessment). The density of microporous PCL determined by weighing 2 mm thick discs was 0.2970.002 g/ cm3. Examination of the material by SEM revealed a surface comprising flat, smooth areas and fibrous regions. The underlying morphology was produced by an array of crystalline, lamellar structures giving rise to pores of irregular size and shape of the order of 1–10 mm dimensions (Fig. 1A). No changes in morphology were evident after incubation of samples for 45 months in PBS at 37 C.
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Fig. 1. Scanning electron micrographs: (A) Microporous PCL; (B) HA/PCL material and (C) Inulin/PCL material.
The density of HA/PCL material was 0.4970.003 g/ cm3. SEM examination revealed the irregular, porous morphology formed by the PCL lamellar structures (Fig. 1B). HA particulates less than 5 mm in size were evident, dispersed and embedded within the PCL phase. The density of inulin/PCL material was 0.8970.002 g/ cm3. SEM examination revealed agglomerates of roughly spherical inulin particles, approximately 5 mm in size (arrowed in Fig. 1C) enmeshed in a PCL matrix. No changes in morphology were evident after incubation of samples for 45 months in PBS at 37 C. 3.1. Weight loss characteristics Weight loss was not recorded for microporous PCL samples in PBS at 37 C over the 45 month time course of the study. This is expected on the basis of past work which revealed that significant weight loss of PCL implants does not occur until the molecular weight falls to around 5000 [21]. The inulin/PCL materials exhibited a mean weight loss of 10.2% after 1 month in PBS and 17.6% after 12 months. No further weight loss was recorded up to 45 months.
3.2. Molecular weight changes The molecular weight characteristics of microporous PCL and inulin/PCL materials pre- and post-incubation in PBS at 37 C are presented in Table 1. Relatively minor changes were measured in molecular weight and differences in molecular weight, respectively, between each sample type over 12 months, reflecting the slow degradation rate of PCL. At extended degradation times of 32 and 45 months, microporous PCL tends to exhibit a more rapid decrease in molecular weight than the inulin/PCL possibly due to a higher porosity of the unloaded material (Fig. 1) which both facilitates permeation of the PBS incubation medium and increases the material surface area for hydrolysis. 3.3. Thermal analysis The melting point and Tg of microporous PCL and inulin/PCL materials were, in general, close to the values usually quoted for PCL i.e. 60 C and 60 C, respectively (Table 2). The higher value of Tg recorded for microporous PCL incubated for 45 months in PBS
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Table 1 The molecular weight characteristics of microporous PCL and inulin/ PCL materials pre- and post-incubation in PBS at 37 C Sample
Incubation time (months)
Mn
Mw
Polydispersity
PCL
0 1 2 3 6 12 32 45
78,778 77,999 77,406 75,040 75,550 74,112 55,679 40,137
113,508 112,887 112,090 109,153 110,442 108,011 87,255 79,710
1.44 1.45 1.45 1.46 1.46 1.46 1.57 1.99
0
76,234
110,610
1.45
1 2 4 9 12 32 45
71,713 72,653 75,291 69,164 68,274 60,397 57,182
105,985 105,349 108,878 105,616 102,708 99,933 94,831
1.48 1.45 1.45 1.50 1.50 1.66 1.66
Inulin/ PCL
Table 2 The thermal characteristics of microporous PCL and inulin/PCL materials pre- and post-incubation in PBS at 37 C Tg ( C)
Sample
Incubation time (months)
Tm ( C)
PCL
0 6 12 32 45
60.3 60.6 62.5 61.4 61.8
61.6 63.5 62.0 59.2 55.9
77.9 78.0 78.6 77.4 79.8
Inulin/PCL
0 4 12 30 44
61.8 62.7 62.2 62.6 61.9
59.0 61.7 58.8 58.3 57.9
61.6 62.0 55.9 56.2 54.2
55.3
62.8
59.0
Unprocessed PCL pellets
Crystallinity (%)
suggests a constraint on chain mobility in the amorphous phase. High crystallinity of approximately 80% developed in microporous PCL during production and remained unchanged during incubation in PBS at 37 C. In contrast the crystallinity of inulin/PCL materials was around 25% lower (similar to unprocessed pellets) and also remained unchanged during incubation in PBS. 3.4. Indentation resistance Microporous PCL materials are resilient under small strains but can be compressed readily resulting in permanent deformation. Values of the indentation
Table 3 The indentation resistance of microporous PCL and inulin/PCL materials following incubation in PBS at 37 C Immersion time in PBS (months)
0 1 2 3 4 6 9 12 32 45
Indentation resistance (mn/m2) PCL
Inulin/PCL
1.270.3 1.270.2 0.970.5 1.170.5 ND 1.170.1 ND 1.170.03 1.070.1 1.270.1
24.076.0 22.172.1 10.970.5 ND 10.470.8 ND 11.270.7 11.673.6 9.674.0 10.070.2
resistance of microporous PCL and inulin/PCL materials, pre- and post-immersion in PBS at 37 C are listed in Table 3. The indentation resistance of microporous PCL was virtually unchanged after 45 months immersion in PBS at 37 C at around 1.2 MN/m2. The force/deflection curve also remained unchanged by immersion of samples in PBS, consisting of a gradual, almost linear increase in load with increasing material compaction and density below the indentor probe (Fig. 2D). The minor changes recorded in indentation behaviour correlate with the hydrolytic stability of PCL as indicated by molecular weight determinations (Table 1). Although differences in test methods preclude strict comparisons, the indentation resistance of microporous PCL is similar to the compressive strength (1.0 MN/m2) of microporous PLA/PLG blends prepared by gel casting [23]. The inulin/PCL materials are hard and rigid materials due to the high (67% w/w) loading of inulin powder. This gives rise to an indentation resistance which is almost 20 times higher than microporous PCL (Table 3). In contrast to microporous PCL a major decrease in compression resistance of almost 70% was measured for inulin/PCL materials after 2 months immersion in PBS at 37 C which then remained fairly constant over the next 43 months. The reduction in mechanical properties of the inulin/PCL materials coincides with a loss in sample weight of 11% due to dissolution and extraction of the reinforcing polysaccharide phase. Distinct changes in the indentation force/deflection curve were observed with time of incubation in PBS. Up to 12 months, a two-stage curve was apparent exhibiting a yield point followed by an almost linear load/deflection trace of lower slope (Fig. 2A). This behaviour is similar to that reported for microporous gel-cast materials produced from PLA, PLG and their blends [23]. After 12 months, inulin/PCL materials gave rise to a yield point but fractured before the test limit of deflection
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resulting in radial crack propagation and a gradual decrease in load-bearing ability (Fig. 2B). An increase in incubation time to 32 months and above resulted in fracture of inulin/PCL samples without yielding, gradual crack propagation and decreased load-bearing ability (Fig. 2C). 3.5. Osteoblast-interaction with PCL and PCL/HA materials
extraction may be applicable such as PLA or high llactide-containing PLG copolymers. Various materials may be added to the PCL solution to adjust the properties of the finished material such as degradation rate, morphology and the pattern of drug release, or to modulate cell interaction with the material. The absence of any weight loss for microporous PCL materials is in line with the findings of Pitt et al. [21] who investigated the degradation of implanted, sub-dermal, drug delivery devices produced from PCL (Mn 50,000) over a timescale of 52 months. Weight loss was not recorded until after approximately 2.5 years when the molecular weight (Mn ) had fallen to 5000, reflecting the requirement for degraded chain fragments to be below a critical chain length for diffusion from the matrix. PCL is a semi-crystalline aliphatic polyester. As such the degree of crystallinity is important in controlling degradation since the crystalline regions retard fluid ingress. The method of manufacture of microporous
60 A
40 B
Force (N)
The characteristic HOB cell morphology observed on the various substrates after 90 min, 4 and 24 h in cell culture is indicated in Table 4 and illustrated in Figs. 3 and 4. The degree of cell spreading was higher on Thermanox disks than on PCL films at time points of 90 min and 4 h and equivalent at 24 h. The notable differences in cell morphology on microporous PCL and HA/PCL materials compared with PCL films and Thermanox relates to the coalescence of attached cells after 4 h in culture. Individual cells were still discernible on Thermanox up to 24 h (Fig. 3). In contrast, Fig. 4D shows HOB cell layers draped over the surface features of HA/PCL material after 4 h, masking the underlying grainy texture. The tops of rounded cells assist in identifying the areas of cell coalescence. In Fig. 4B a sheet of HOB cells formed after 24 h in culture spans the underlying fibrous surface structure of microporous PCL.
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4. Discussion Blocks of microporous materials have been formed by gradual precipitation from solutions of PCL in acetone induced by solvent extraction across a semi-permeable membrane which is formed in situ at a PCL solution/ methanol interface. The viscosity of PCL solutions above 10% w/v concentration is sufficiently high to resist mixing with the non-solvent phase. This behaviour coupled with rapid production of an interfacial film results in the separation of two distinct fluid layers. Solvent exchange across the film subsequently induces gradual precipitation of PCL to form a block of microporous material. Although PCL provided the focus for the present work, other polymers which show a tendency to crystallise from solution on solvent
C 20
D 0.5
0
1
Deflection (mm) Fig. 2. Characteristic indentation force–deflection curves for inulin/ PCL materials (A,B,C) and microporous PCL following incubation in PBS at 37 C: (A) inulin/PCL, 0–12 months in PBS; (B) inulin/PCL, 12–32 months in PBS; (C) inulin/PCL, >32 months in PBS and (D) microporous PCL.
Table 4 The morphology of primary human osteoblasts cultured on tissue culture plastic (Thermanox), PCL films and microporous PCL materials Time point
90 min 4h 24 h
Appearance of HOB cells Microporous PCL and HA/PCL
PCL film
Thermanox
Rounded, cell processes visible Spread, forming cell layer Cell layer, individual cells not distinguishable
Rounded Rounded and spread cells Spread cells
Rounded and spread cells Spread cells Individual cells discernible
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(A)
(B)
RC
SC
(C)
SC
Fig. 3. The morphology of human osteoblasts following cell culture on PCL films and Thermanox. (A) Rounded cells (RC) on PCL film after 90 min in cell culture; (B) Rounded and spread cells (SC) on Thermanox after 90 min and (C) Spread cells on Thermanox after 4 h in cell culture.
PCL involving slow, gradual precipitation from solution allows the development of high levels of crystallinity similar to the behaviour of solvent-cast PLA films [23]. This essentially restricts the scope for further crystallization during incubation in PBS (Table 2). In contrast Pitt et al. [21] measured a rapid increase in crystallinity of PCL drug delivery devices over the first 4 weeks postimplantation from 45% to 50% which was ascribed to annealing of the polymer at body temperature. The crystallinity gradually increased to a value of approximately 80% after 30 months which was attributed to crystallization of tie chain segments from the amorphous phase following chain cleavage, facilitated by the low Tg of PCL. The relative insensitivity of the crystallinity of inulin/PCL materials to incubation in PBS in the present study (Table 2) suggests that the PCL chain segments in the amorphous phase are strongly bound to the polysaccharide phase thereby restricting the molecular mobility necessary for reorganization and crystallization. The constancy of mechanical performance of microporous PCL samples following immersion in PBS at 37 C for 45 months suggests that the reduction in fracture toughness observed for inulin/PCL materials at
and beyond 12 months is due to changes in the structural stability of inulin. This behaviour may be exploited to facilitate the breakdown of inulin-loaded implants in vivo by tissue ingrowth, thereby enhancing the rate of elimination of implanted material by phagocytosis. Osteoblasts display distinct morphological changes on biomaterials, often within 2 h of cell culture in the case of HA and alumina surfaces [24]. Cytoskeletal reorganisation (cell flattening/spreading and process extension) is anticipated by 4 h following contact of rounded cells with the surface while cells are expected to be flat, spread and to have started proliferating after 24 h [25]. The timing of the above morphological changes is not exact and varies with substrate characteristics and cell type [16,25,26]. Spreading of HOB cells occurred on microporous PCL and HA/PCL materials in the same time period as Thermanox which is surface modified to promote favourable cell attachment and growth characteristics. HOB cells were found to be spread to a larger extent on microporous PCL and HA/PCL materials at 4 h than on PCL films. These differences may be partly explained in terms of competitive and/or selective adsorption of cell adhesion molecules, such as fibronec-
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(B)
(A)
P
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P CL D CL
(D)
(C)
CL CL
CL
RC
Fig. 4. The morphology of human osteoblasts following cell culture on microporous PCL and HA/PCL materials. (A) Microporous PCL after 90 min in cell culture. (B) Microporous PCL after 24 h. CL cell layer, P-polymer, D-identifying dehydration cracks. (C) HA/PCL material after 90 min in cell culture. (D) HA/PCL after 4 h.
tin, to produce a favourable packing density and conformation-controlled presentation of adhesion sequences. HA is known to enhance cell adhesion and spreading probably through an inherently high capacity for adsorbing proteins [27]. In addition, the surface topography of the microporous PCL materials may be more conducive to cell attachment and development than a smooth film surface. The numbers of adhered cells, cell shape and attachment forces have been shown by many investigators [13,18] to be affected by surface morphology. The propensity for cell coalescence to form sheets on the microporous PCL and HA/PCL materials (Fig. 4) rather than separated cells indicates the presence of a high density of anchorage sites for integrin binding leading to extensive spreading and cell–cell contacts. This in turn may positively influence the pattern of mechanical stresses experienced by the cytoskeleton [28]. Cell coalescence on PCL and HA/PCL materials may also favour formation of the osteoblast sheets considered advantageous for correct development of bone tissue [29]. The low resorption rate of precipitation cast materials based on PCL recommends their use as long-term,
delivery systems for bioactive molecules such as growth factors and hormones and extended-residence supports for cell growth and tissue development. Growth factors such as BMP may be incorporated in PCL microporous materials by simple admixing of spray-dried particulates, thereby providing a bone graft substitute which both supports and stimulates local bone growth to achieve integration with host tissue. The potential for controlling drug release from PCL matrices produced by precipitation casting was evaluated using inulin polysaccharide (Mw 5000) as a model biopharmaceutical. Inulin has been used to investigate the release of macromolecules from collagen gels [30] since inulin is filtered by the kidneys and excreted rather than being bound to plasma proteins or metabolized. Inulin release from a drug delivery device in vivo may thus be directly related to the excretion rate. Although the duration of inulin release from PCL matrices was extended to 12 months, the cumulative release was limited to approximately 18% w/w, indicating that the polysaccharide phase is efficiently coated by PCL and pore connectivity is poorly developed. Water soluble polymers are frequently incorporated in polymeric films and
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microspheres to modulate drug delivery [31]. This approach may be useful for improving the rate and duration of macromolecule release from microporous PCL matrices. Copolymerisation and polymer blending provide alternative approaches via control of the resorption characteristics of the matrix.
5. Conclusions Precipitation casting presents a novel approach for producing microporous polymeric, block-form materials and particulate-loaded compositions. The low resorption rate recommends application of PCLbased castings for long-term, release of bioactive molecules and extended-residence cell supports in tissue engineering.
Acknowledgements Thanks are due to Tony Evans, Solvay Interox, UK for providing the PCL molecular weight data in Table 1. AGAC acknowledges the prior financial support of the Division of Orthopaedics and Accident Surgery, Queen’s Medical Centre, University of Nottingham.
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