J Mater Sci: Mater Med (2013) 24:1217–1227 DOI 10.1007/s10856-013-4881-x
Examining porous bio-active glass as a potential osteo-odonto-keratoprosthetic skirt material Reeta Huhtinen • Susan Sandeman • Susanna Rose • Elsie Fok • Carol Howell Linda Fro¨berg • Niko Moritz • Leena Hupa • Andrew Lloyd
•
Received: 8 October 2012 / Accepted: 28 January 2013 / Published online: 6 February 2013 Ó Springer Science+Business Media New York 2013
Abstract Bio-active glass has been developed for use as a bone substitute with strong osteo-inductive capacity and the ability to form strong bonds with soft and hard tissue. The ability of this material to enhance tissue in-growth suggests its potential use as a substitute for the dental laminate of an osteo-odonto-keratoprosthesis. A preliminary in vitro investigation of porous bio-active glass as an OOKP skirt material was carried out. Porous glass structures were manufactured from bio-active glasses 1-98 and 28-04 containing varying oxide formulation (1-98, 28-04) and particle size range (250–315 lm for 1-98 and 28-04a, 315–500 lm for 28-04b). Dissolution of the porous glass structure and its effect on pH was measured. Structural 2D and 3D analysis of porous structures were performed. Cell culture experiments were carried out to study keratocyte adhesion and the inflammatory response induced by the porous glass materials. The dissolution results suggested that the porous structure made out of 1-98 dissolves faster than the structures made from glass 28-04. pH experiments showed that the dissolution of the porous glass increased the pH of the surrounding solution. The cell
R. Huhtinen N. Moritz BioCity Turku Biomaterials Research Program, Institute of Dentistry, Turku Clinical Biomaterials Centre (TCBC), University of Turku, Turku, Finland S. Sandeman (&) S. Rose E. Fok C. Howell A. Lloyd Biomaterials and Medical Devices Research Group, School of Pharmacy and Biomolecular Sciences, University of Brighton, Huxley Building, Lewes Road, Brighton, East Sussex BN2 4GJ, UK e-mail:
[email protected] L. Fro¨berg L. Hupa ˚ bo Akademi University, Process Chemistry Centre, A Turku, Finland
culture results showed that keratocytes adhered onto the surface of each of the porous glass structures, but cell adhesion and spreading was greatest for the 98a bio-glass. Cytokine production by all porous glass samples was similar to that of the negative control indicating that the glasses do not induce a cytokine driven inflammatory response. Cell culture results support the potential use of synthetic porous bio-glass as an OOKP skirt material in terms of limited inflammatory potential and capacity to induce and support tissue ingrowth.
1 Introduction A keratoprosthesis (KPro) is a corneal implant which is used in the eye to replace the central area of an opacified cornea in patients who are unresponsive to donor corneal transplant. Typically, the keratoprosthesis comprises a transparent part which is capable of transmitting light from the exterior of the eye to the retina. The optical central part is surrounded by a supporting skirt which keeps the keratoprosthesis anchored within the cornea. The early keratoprostheses were fabricated using glass or quartz glass to provide a convex disc in the central part of the keratoprosthesis, allowing light to travel to the back of the eye. In the early designs keratoprostheses were made totally out of glass or alternatively the central part was made from glass and metal rings or flanges were used to anchor the keratoprothesis within the eye. Supporting structures for glass were made out of metals like gold, platinum, tantalum and stainless steel. One problem with the early KPro designs was the loss of tissue around the prosthetic rim. In 1862, Abbate used a keratoprosthesis consisting of a glass disk surrounded by a skirt made of two successive rings: guttapercha and casein, both natural polymers [1]. Although the
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choice of peripheral prosthetic materials was not very successful, the approach indicated the need for a skirt material to promote better incorporation of the prosthesis into the host cornea. In 1937 Salzer made two revolutionary suggestions. Firstly, keratoprostheses should be made of materials lighter than glass and, secondly, the prosthetic skirt should be made of materials which are able to bond tightly with the host tissue [1]. The modern period in the development of keratoprosthesis began with the use of poly(methyl methacrylate) (PMMA). PMMA is transparent and lighter than glass. Typical density values are 1.1 for PMMA and 2.5 g/cm3 for glass. At present, PMMA is one of the most commonly used core materials in so called ‘‘core and skirt’’ KPro design where the core consists of an optically transparent material while the skirt forms the supporting structure for the optical part. Several other polymers have been tested in KPro designs with varying success. These include silicone (Aachen-KPro) [2], poly(2-hydroxyethyl methacrylate) PHEMA (AlphaCorÒ previously known as Chirila KPro) [3], polyurethanes (Seoul KPro) [4, 5], polytetrafluoroethylene (PTFE) TeflonÒ [6, 7], polyethylene terephthalate DacronÒ [8], a fibrous melt-blown web of polybutylene and polypropylene [9, 10] and carbon fibres [11]. Teflon is no longer considered a suitable material in corneal surgery due to several shortcomings; it has poor adhesion to the surrounding tissue and, especially if used in porous form, it provides a potential route for infection. Theoretically, the KPro should integrate with the corneal epithelium, stroma, endothelium, or a combination of these corneal layers. Integration of the KPro with the corneal epithelium, which is the outmost layer of cornea, may act as a barrier to infections, but offers little structural support. Stromal adhesion has been reported to increase the structural integrity of a KPro [12]. The Stroma comprises about 90 % of the cornea thickness and it contains mainly collagen fibres, glycosaminoglycans and keratocytes. The advantage of a porous skirt material is that it could allow stromal keratocytes to penetrate, proliferate, and synthesize connective tissue proteins inside the skirt structure. When this happens, ‘‘healing’’ will occur, creating a natural anchor between the synthetic material and host tissue. Additionally, there are no blood vessels in the cornea and the porous structure would also allow the transport of nutrients into the eye even if the material itself is nonpermeable. There have been various attempts to enhance the bond between the KPro and surrounding tissues through the promotion of better cell adhesion. In order to promote epithelial cell adhesion the surface has been modified with naturally occurring extracellular matrix proteins like collagen [13, 14], fibronectin [15, 16] and laminin [17]. Another way to increase the biocompatibility of the
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prosthesis is to use autologous tissue, like tooth and bone, as a skirt material around a PMMA core. In Strampelli’s osteo-odontokeratoprosthesis (OOKP) [18] the patient’s own tooth and part of the jaw bone are used to form a biocompatible skirt around the PMMA core. A relatively good, long term clinical outcome can be achieved with the OOKP, but the disadvantages are that the OOKP surgery has to be done in several steps, it is time consuming and the patient’s own tissues have to be sacrificed. It has been asserted that as long as the grafted tissue stays viable in the eye the prosthesis will stay in the right position. As both tooth and bone are porous they enhance tissue ingrowth into the skirt matrix and thus increase the integration of the OOKP in the eye. However, a long lasting inflammation can locally decrease the pH of the tissue and cause the degradation of tooth and bone leading to loosening of the prosthesis and finally the loss of the OOKP [19]. After the early development of KPro materials, so-called bio-active glasses have been developed and tested in vitro and in vivo. They are synthetic, silica derived materials, which are able to bond with living tissues. Bio-active glasses also have osteoinductive capacity. Currently bio-active glasses are used clinically as fillers for bone cement and in dental restorative composites mainly in the form of granulates or plates. Certain glass compositions can also be manufactured into structures with interconnected porosity [20]. In bone replacement applications the bio-active glass slowly resorbs and is finally replaced by newly formed tissue. The influence of glass composition on in vitro reactions was recently reported [21]. Bioactive glasses are interesting material options for OOKPs, as they bond with living tissue. However, one problem with typical bio-active glasses is that they resorb with time, thus limiting their use in OOKP design. It is known that the dissolution of glasses in neutral and acidic solutions commences by an ion exchange of alkalis and alkaline earths ions from the glass to hydrogen in the solution. This reaction increases the pH of the immersion solution and also gives a silica rich layer on the glass surface. However, the in vivo dissolution of glass showing bioactivity is not properly established. Soft tissue bonding has been reported only for glasses which show a high degree of bioactivity, and thus a relatively rapid resorption [22]. Bioactive glass–ceramic coating on titanium in an ocular environment in rabbits was found to cause problems due to degradation and detachment of the coating. The detachment was assumed to be partially due to coating thickness [23]. Some bio-active glass compositions have been reported that react slowly in vitro [21]. Among these, several compositions allow the manufacture of porous implants without interfering with simultaneous crystallization of the glass [24]. These compositions could be of interest in OOKP skirt design. This study investigated the suitability of porous bioactive glasses as replacement OOKP skirt materials. The most suitable glass would be a relatively stable base
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material which promotes peripheral apatite deposition to allow bonding to surrounding tissue without the danger of widespread implant erosion and extrusion of the implant by newly forming tissue. An interwoven backbone material may be necessary to hold the OOKP optic in place if a slowly resorbing bio-glass is found to be the most suitable option. Skirt structures with interconnected porosity were manufactured of two glass compositions and tested in an in vitro environment. Degradation of the porous structures was studied in simulated aqueous humour and also changes in pH were measured. 2D porosity measurements are based on SEM-image analysis and 3D structural analysis on micro-computed tomography. The interaction of keratocytes with the porous glass was investigated in order to assess the possibility that bio-active glass may induce corneal in-growth and integration without exacerbating corneal inflammation and, in this respect, be suitable as an OOKP skirt material. Inflammatory response to the glass samples was measured by cell expression of the cytokines IL-6 and IL-8 which were found to be expressed by cultured human keratocytes and are known to modulate the corneal response to inflammation.
2 Materials and methods 2.1 Preparation of porous skirts Two glass types 1-98 and 28-04, which according to in vitro experiments using simulated body fluid (SBF), show medium and low level bioactivity respectively, were chosen [21]. Both glass compositions allow a viscous sintering of particles into porous bodies without crystallization [25]. The oxide composition of the glasses is presented in Table 1. The glasses were produced by melting of commercial grade Belgian sand and analytical grade Na2CO3, K2CO3, MgO, CaCO3, H3BO3 and CaHPO42H2O in a Pt crucible at 1360 °C for 3 h. After casting and annealing, the glasses were remelted to improve homogeneity. The glasses were crushed and sieved into particles within the size ranges 250–315 and 315–500 lm. The particles were sintered in a graphite mould into ring shaped structures with interconnected porosity in an electrical furnace in a nitrogen atmosphere at 710 °C. The structures were given the codes 98a, 04a and 04b according to Table 2. The porosity of the skirt structures was adjusted by using two different sintering times 60 and 90 min. The sintering time was calculated from
the moment the sample reached the preset temperature, i.e. roughly 5 min after inserting the sample into the furnace. After sintering, the porous structures were cooled in flowing nitrogen. 2.2 In vitro dissolution studies Dissolution of ions from the porous glass structures was measured in simulated aqueous humour. Simulated aqueous humour mimics the inorganic ion composition of human aqueous humour [26]. The ion concentration of the human and simulated aqueous humour is presented in Table 3 [27]. Zinc is present in the human aqueous humour, but in simulated aqueous humour zinc precipitates and is thus excluded. The following chemicals were used to prepare the simulated aqueous humour: sodium chloride 99.5 % (Sigma), potassium chloride ACS reagent (Sigma), sodium bicarbonate 99.5 % (Sigma), potassium phosphate dibasic trihydrate 99 % (Sigma) and tris(hydroxymethyl)aminomethane 99.8 % (Aldrich). Tris buffer was used to adjust the pH to 7.4. Dissolution experiments were done at 37 °C in a shaking water bath at 150 rpm (Heto, Heto Lab Equipments, Denmark). Silica and calcium ion concentrations in the simulated aqueous humour were measured colorimetrically. Calcium ion concentration measurement is based on an ortho-cresolphthalein complex method and silica measurement on the molybdenum blue complex method. Absorbances were measured using a UV-1601 spectrophotometer (Shimadzu, Australia) and Multiskan MS ELISA plate reader (Labsystems, Finland) respectively. The measured silica and calcium ion values were used to calculate the dissolution of silica and CaO from the glass. The weight of each porous glass sample varied between 0.99–1.20 g and the volume of simulated aqueous humour was 100 ml per sample in the beginning of the in vitro dissolution experiment. Part of the solution was replaced during the experiment to avoid saturation and absolute Ca2? and silica concentrations at each time points were measured. The cumulative release of silica and calcium oxide is calculated and results are presented as percent of loss of the nominal glass composition. The measured values are given as average values of three samples. 2.3 Change in the pH of simulated aqueous humour Although the aqueous humor is a buffered solution, dissolution of bio-active glass is likely to increase its pH.
Table 1 Nominal oxide composition in wt% (mol%) Glass
Na2O
K2O
MgO
CaO
B2O3
P2O5
SiO2
1-98 28-04
6 (5.9) 5 (4.9)
11 (7.1) 11.25 (7.2)
5 (7.6) 6 (9.0)
22 (23.9) 15 (16.2)
1 (0.9) 3 (2.6)
2 (0.9) 0 (0)
53 (53.8) 59.75 (60.1)
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Table 2 Particle size fractions and sintering parameters of porous glass skirts Material code
Glass
Particle size (lm)
Sintering temperature (°C)
Sintering time (min)
98a
1-98
250–315
710
60
04a
28-04
250–315
710
60
04b
28-04
315–500
710
90
Table 3 Ion concentration of human aqueous humour and simulated aqueous humour Element
Human aqueous humor (mmol/l)
Simulated aqueous humor (mmol/l)
Sodium (Na?)
111
125
Potassium (K?)
3.2
3.2
Chloride (Cl-) Bicarbonate (HCO3-)
107 20.6
107 20
Phosphorus (P)
0.62
0.62
Zinc (Zn)
1.59
–
Tris
–
50
The pH of the solution during immersion of the porous glass structures was measured at different time points. The glass to aqueous humour ratio was 20 mg/ml. The immersion solution was not changed during the experiment. The experiment was carried out at 37 °C in a shaking water bath at 150 rpm (Heto, Heto Lab Equipments, Denmark). Simulated aqueous humour without porous glass was used as a control solution and the pH of the control solution was measured during the experiment. The pH was measured with a PHM220 Lab pH Meter (Radiometer) using a pHC2401 combination pH electrode. The solution temperature was measured with a T201 thermometer (Radiometer). Prior to experiments the pH electrode was calibrated with Radiometer analytical buffer solutions at pH 7.00 and pH 10.00. 2.4 Structural analysis of porous samples 2D porosity of the glass structures 98a, 04a and 04b was measured by image analysis of SEM micrographs of the sample cross-section. Three to six samples of each glass structure were cast in EpoFix Kit (Struers) and cut in order to reveal the cross-sections. The cross-sectional surfaces were polished. Scanning electron microscope (SEM) measurements were obtained with a Leo Gemini 1530 SEM using a Thermo NORAN Vantage X-ray detector. Adobe Photoshop Elements 6.0- and ImageJ- programs
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were used to edit the SEM-images into black and white images, which were used to analyse the relative 2D pore/ glass surface area ratio. 3D structural analysis was done using micro-computed tomography (micro-CT, Skyscan 1072, Skyscan n.v. Kontich, Belgium) for porous structures 98a, 04a and 04b. To obtain reasonable resolution with micro-CT imaging, a sector of around 90° was cut away from each of the original ring-shaped samples. The samples were cut using a highspeed diamond disc attached to a dental drill. Thereafter, each sample was mounted on the standard sample holder and imaged separately. The specimens were imaged with the step angle of 0.675° within the full angle of 180°. Source voltage was 53 kV, source current was 189 lA, and no filters were used. In image acquisition, a single 16 bit grayscale shadow projection image was obtained for each step angle, whilst no frame averaging was used. All imaging and reconstruction parameters were kept identical for each of the specimens imaged. For each specimen, the acquired shadow projection images were reconstructed into an array of cross-sectional 8 bit grayscale images using NRecon software (Skyscan n.v. Kontich, Belgium). Automatic post-alignment and beam hardening correction were used in the reconstruction. The resulting spatial resolution of the cross-sectional images was 5.86 lm per voxel. Arrays of cross-sectional images were further used to study structural features of the specimens. 3D structural analysis of the three-dimensional data arrays was performed using CTAn software (Skyscan n.v. Kontich, Belgium). As the first step of the analysis, the contour of the specimen was manually traced on the top and bottom cross-sectional images in the image arrays. The three-dimensional volumes of interest (VOI) obtained consequently served as bounding VOIs in the analysis. The volume of these bounding VOIs served as a measure of the total volume (TV) of the specimen. Material volume (BV) was established as a result of global thresholding. Thereafter, the software provided direct calculations of morphometric parameters including material volume fraction (BV/TV), mean trabecular thickness (TbTh), trabecular separation (Tb.Sp), and trabecular number (TbN). Nonmetric parameters, such as trabecular bone pattern factor (Tb.Pf), structure model index (SMI) were also calculated. 3D porosity was calculated as (TV-BV)/TV expressed as percentage and presented in the results section together with mean pore sizes. It is also possible to calculate the material surface area (BS) per material volume as BS/BV. The analysis software also gave the distribution of trabecular thicknesses (distribution of distances between the pores) and the distribution of trabecular separations (distribution of pore sizes), both are presented in the results section.
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2.5 In vitro assessment of keratocyte adhesion and cytokine production
content according to manufacturer’s instructions (BD Biosciences). Media control (without cells) was also assessed to determine any background levels of cytokine.
2.5.1 Cell culture and characterisation of cytokine expression profile
2.5.2 Cell adhesion and cytokine production in response to bio-glass exposure
A human keratocyte cell strain established at the University of Brighton from donated corneal tissue to the East Grinstead Eye Bank was used for the in vitro cell studies. Cells were grown in DMEM supplemented with 10 % foetal calf serum and 1 % penicillin/streptomycin and were passaged by standard trypsinisation using 19 trypsin/EDTA. Cells at both early and mid-phase cumulative population doublings were assessed for the expression of vimentin and cytokeratin by immunocytochemistry to determine maintenance of fibroblast phenotype in culture. The secretion of cytokines IL-1b, TNF, IL-8 and IL-6 were measured by enzyme linked immunosorbent assay (ELISA). Keratocytes were seeded into the wells of a 24 well plate at a concentration of 1 9 105 cells in 1 ml of media per well. Wells were set up for each time point with either no LPS, 10 ng, 100 ng or 1,000 ng/ml LPS added. Samples of cell conditioned media were collected after 1, 2, 6, 24, 30 and 48 h. After centrifugation at 3009g for 5 min to remove debris, samples were frozen at -80 °C 96 well ELISA plates were coated in capture antibody diluted 1:250 in carbonate buffer and incubated at 4 Æ C overnight. Samples were diluted in assay diluent (10 % FCS in PBS) at a range of different concentrations from 1:2 to 1:200 and assayed for cytokine
Three of each porous glass structures 98a, 04a and 04b were placed in a 24 well plate and incubated in 1 ml of supplemented DMEM (10 % FCS, 1 % penicillin/streptomycin) for 24 h at 37 °C. The media was removed and the samples were washed in PBS three times. A human keratocyte cell strain was passaged and 1 9 105 cells suspended in 100 ll supplemented DMEM was added to each sample. Cells were added in clockwise fashion directly onto the surface of the samples in 10 amounts of 10 ll. Bacterial lipopolysaccharide (LPS) positive and LPS negative controls were set up in triplicate. For the controls 100 ll cell suspension was added directly to the tissue culture well and 0.9 ml of media was added immediately to prevent cell drying. The glass samples were incubated for 1 h to allow cell adhesion. 0.9 ml of media was added to each of the sample wells. 10 ll of a 1 lg/ml LPS solution was added to 1 ml of media in the LPS positive control wells to give 10 ng/ml LPS final concentration. The plates were incubated for a further 5 h. Media was removed and stored at -20 °C for measurement of cytokine production by ELISA. A further 1 ml of media was added to the wells and plates were incubated overnight. Media was removed and wells were rinsed with PBS. 1 ml
10
120
98a
a
100
04a 04b
80 60
98a
c
8
04a 04b
6 4
40 2
20
0
0 0
100
200
300
400
500
0
100
60
98a 04a 04b
b
50
200
300
400
500
Time [h]
Time [h]
40
10
98a
d
8
04a 04b
6
30 4
20
2
10
0
0 0
100
200
300
400
500
Time [h]
Fig. 1 a Silica concentration in simulated aqueous humour as a function of immersion time b Ca2? concentration in simulated aqueous humour as a function of immersion time c Cumulative
0
100
200
300
400
500
Time [h]
dissolution of SiO2 from porous glass structures d Cumulative dissolution of CaO from porous glass structures
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7,60
the pH of the pure simulated aqueous humour increased with immersion time. This has been taken into account in the results presented in Fig. 2 as a background subtraction. The increase of the pH in the blank solution was likely due to carbon dioxide evaporation and thus shifts in the carbonate equilibrium value during the pH experiment.
7,50
3.3 Structural analysis
7,80 98a 04a
pH
7,70
04b
7,40 0
100
200
300
400
500
Time [h]
Fig. 2 The pH change caused by the porous glass structures in simulated aqueous humour as a function of immersion time
of a 1 lg/ml calcein-am solution in media was added to each well and plates were incubated at 37 °C for 30 min. Cell adhesion was observed using fluorescence microscopy, counting cell adhesion in a total of 15 fields for each sample in clockwise fashion. Cell penetration into the materials was observed using confocal microscopy following DAPI staining of the cell nuclei. Cell conditioned media was assayed for the presence of cytokines IL-8 and IL-6 according to manufacturer’s instructions (BD Biosciences). Samples underwent a 1:100 dilution in assay diluent.
Fig. 3 shows edited black and white SEM images of the polished cross-sections of 98a, 04a and 04b. The glassy parts can be seen as a white structure interspersed with a black matrix indicating the pores. Average 2D porosity was calculated from edited SEM-images as the relative surface area of pores versus the total area of pores and glass (Table 4). In the Table 4 is also given the average 3D porosity and the average pore sizes as well as the surface area to material volume ratios calculated from the values obtained with micro-computed tomography analysis. Although the absolute values are smaller for the 2D porosity than the 3D porosity for each sample the porosity increased in the same order: 98a \ 04b \ 04a. Where as the 3D average pore size increased in the following order: 98a \ 04a \ 04b. The average distance between pores is longer in 04b then in 98a and 04a as clearly seen also in the SEM-images. Distribution of pore sizes and distances between pores are presented in Figs. 4 and 5 respectively.
3 Results
3.4 Keratocyte adhesion and inflammatory response
3.1 In vitro dissolution
The keratocyte cell strain reached 35 cumulative population doublings (cpd) prior to senescence of the cultures. Cells at early and mid-phase cpd showed positive expression of vimentin and negative expression of cytokeratin. Since cytokeratin is found in corneal epithelial cells but not in stromal keratocytes and vimentin is present in cells of mesenchymal origin the results confirmed that the cultures were uncontaminated keratocytes. No IL-1b or TNF cytokine secretion by the keratocytes was measured over a period of 48 h in the absence and presence of LPS. Significant IL-6 and IL-8 production occurred at the 6 h time point and reached a plateau at 24 h. IL-6 and IL-8 secretion was much greater in the presence of LPS (Fig. 6). However, increasing the LPS concentration from 10 to 10,000 ng/ml did not increase the concentration of IL-6 or IL-8 secreted any further. Therefore an LPS concentration of 10 ng/ml was used for further studies. Keratocyte adhesion was seen on the surface of each of the bio-active glass samples. Cells were generally spread out in typical fibroblast type morphology. Surface cell adhesion was greatest for porous glass structure 98a (Fig. 7). A typical field of view is shown for each of the porous structures in Fig. 8. Confocal scans carried out on calcein-am and DAPI
Concentrations of silica and calcium ions in the simulated aqueous humour increased as a function of immersion time (Fig. 1a, b). Amorphous silica has a solubility of 100– 140 ppm in water. Thus Fig. 1a suggests that for all samples the dissolution of the glasses took place under saturated conditions, but not in sink conditions. Structure 98a showed the highest SiO2 and Ca2? concentrations at all time points. The calculated cumulative dissolution of CaO and SiO2 from the porous glass structures is shown in Fig. 1c, d. The relative dissolution of CaO from all structures was higher than the relative dissolution of SiO2, especially in the case of the structure 98a. The dissolution of porous structures decreased in the following order: 98a [ 04a [ 04b. 3.2 Change in the pH of simulated aqueous humour The pH change caused by immersion of the glass structures in simulated aqueous humour is presented in Fig. 2. During the first 100 h 98a gave the highest increase in pH (0.27 pH-units) followed by 04a (0,25) and 04b (0.20). After 100 h the highest increase in pH was caused by 04a. Also
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1223 14 98a
12
04a 04b
10 8 6 4 2 0 0,00
50
100
150
200
250
300
350
400
450
Pore size (µm)
Fig. 4 Distribution of pore sizes measured with micro-computed tomography 9 98a
8
04a 04b
7 6 5 4 3 2 1 0 0,00
100
200
300
400
500
600
Trabecular thickness (µm)
Fig. 5 Distribution of distances between pores measured with microcomputed tomography
Fig. 3 Edited SEM images of a 98a, b 04a and c 04b. In the images white represents glass, black is the pore area and grey is background Table 4 Comparison between 2D and 3D porosity values. 2D values are based on SEM-analysis and 3D on mirco-CT analysis Material code
Glass
Particle size of used glass granulas (lm)
2D pore area (%)
Mean 3D pore volume (%)
Mean pore size (lm)
98a
1-98
250–315
6,8
12,0
92,5
04a
28-04
250–315
23,5
32,9
98,9
04b
28-04
315–500
11,7
21,6
135,4
Also the mean pore size based on Mirco CT-measurement is presented in table
stained cells showed the presence of cells within the porous glass matrix as well as at the surface (Fig. 8). Surface cell adhesion counts may be influenced by sample pore size differences and interparticulate spacing as well as differences in the compatibility of the materials themselves to keratocyte adhesion. While cell counts were
taken of cells adherent to the surface it was evident that, for 04a and even more so for 04b, there were many cells present deeper into the matrix of the material. The surface adherent cells appeared to be sitting on islands surrounded by the open structure of the porous glass. In 04b where pore diameter appears greatest and cells cannot span across the surface, these cells may adhere further within the matrix influencing the reduced surface cell counts observed. Cytokine production by each of the bio-active glass structures was similar to that of the negative control indicating that the glasses do not induce a cytokine driven inflammatory response (Fig. 9a, b). Cytokine production on all of the materials matched that of the LPS negative control of cells on TC plastic. The levels of IL-8 and IL-6 produced in response to LPS, used to indicate a cellular inflammatory response, were significantly higher than those of the keratocytes on the materials or on TC plastic.
4 Discussion Finding a synthetic replacement for the dental laminate of the OOKP remains an unsolved problem since it is difficult to mimic the properties of bone in synthetic form.
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35 30
0 ng/mL LPS 10 ng/mL LPS
25
100 ng/mL LPS 1000 ng/mL LPS
20 15 10 5 0 1
2
6
24
30
48
Time in hours
Fig. 6 Secretion of IL-6 over time obtained from cells adherent to tissue culture plastic in the presence of varying concentrations of LPS. Mean of n = 3 ± SD. Insert fluorescent micrograph shows keratocytes stained positive for vimentin confirming a pure fibroblast culture
Fig. 7 Keratocyte adhesion to the porous glass structures was quantified by fluorescent microscopy, counting 15 fields on each material at 9 200 magnification following cell staining with calceinam (mean ± SEM, n = 3)
Bio-glass can be tuned to match the inorganic/mineral phase of bone, forming a hydroxyapatite like surface layer through ion exchange processes which allow strong localised tissue bonding following implantation. In this respect it may be possible to use bio-glass to replace the OOKP skirt. A bio-glass or composite skirt which quickly stimulates cellular ingrowth would limit complications such as epithelial downgrowth which lead to implant extrusion. However, the biological toxicity effects of soluble ionic species release from the chemically degrading
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bio-glass are unknown. The study investigated the impact of three bio-glass formulations on ionic species release, pH change, rate of dissolution, cell adhesion and inflammatory cytokine release. In vitro dissolution results show that the porous structure 98a sintered from glass 1-98 dissolved faster than the porous structures 04a and 04b from glass 28-04. Glass 1-98 contains a higher concentration of network modifiers Na2O and CaO and less silica than 28-04. This leads to a more open silicate structure of glass 1-98 than 28-04. Thus the dissolution of the non-bridging ions is more likely from glass 1-98 than from 28-04 as is also suggested by the higher calcium ion concentration measured for the structure 98a than for the structures 04a and 04b in Fig. 1d. Structures 04a and 04b have the same chemical composition, but different porosity. The difference in the dissolution behaviour of these structures depends mainly on the difference in their surface area. The concentration of the surrounding solution also affects the dissolution rate of the glasses. After the first 100 h of immersion the SiO2 concentration is close to saturation in the case of 98a and this may slow down the dissolution of the glass. On the other hand, the ion concentration changes caused by the dissolving implant in vivo in the vicinity of the implant are rarely known and it depends on, for example, the size of implant, implantation site, implant dissolution rate and liquid flow in the tissue next to implant. It should be pointed out that the in vitro dissolution results give only a trend between the porous glasses. Their in vivo dissolution rates may vary compared to these in vitro results. pH measurements show that the dissolution of the porous structures 98a, 04a and 04b increased the pH of the surrounding solution. Dissolution of alkali and alkaline earth ions from the glass is compensated by hydrogen bonding which causes the increase of the pH. During the first 100 h the biggest increase in pH was caused by 98a, which showed also the highest relative dissolution of Ca2?. After 100 h the highest pH values were measured for 04a. Porous structures 04a and 04b are made from the same bio-active glass 28-04, but their different effect on pH could be explained by the higher material surface per volume ratio in 04a. Glass 1-98 has higher reactivity in vitro than 28-04 [21], which indicates that calcium phosphate precipitates form faster on 98a than 04a and 04b. pH was measured in a system where immersion solution was not changed during the experiment. This could give a saturated solution with increasing immersion time, and thus prevent a further dissolution of the porous glass structures. The stable pH values after 300 h of immersion can be explained partly by saturation of the solution. In the cell culture experiments pH increase caused by the bio-active glass was taken into account by pre-incubating the glass samples for 24 h. After that porous glasses were washed and used in contact with cells. It is
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Internal Fig. 8 Calcein-am fluorescent images of live keratocyte adhesion to a 04a b 04b c 98a porous glass structures. d Confocal microscope z stack image of DAPI stained keratocyte nuclei showing cells on the porous glass surface of glass 04b and cells within the internal pores
known that cells are sensitive to drastic pH changes and in this way the initial pH increase during cell contact is avoided. It is known that in vivo chronic inflammation plays an important role in OOKP skirt degradation. The acutely inflamed tissue can create an acidic environment. Using bioactive glass structures in such an environment can partly counterbalance the pH changes, and thus have an antibacterial effect. Powdered bio-active glasses have been reported to show antibacterial effect against several bacterium types, mainly because the dissolving glasses effectively increase the pH of the surrounding solution [28]. Figure 5 indicates that the rise in pH produced by bio-glass dissolution levels off below a pH of 7.8 indicating that no extreme rise in pH will occur to the detriment of surrounding tissue.
3D porosity values were higher than corresponding 2D values. The pores are not equally distributed in the structures and 3D analysis is likely to give a more accurate estimation of the porosity. The porous structure 04a showed the highest porosity followed by 04b and 98a, where as 04b had the biggest average pore size. The chemical composition of glass, particle size fraction and sintering time and temperature control the final porosity of the sintered glass structures. In the manufacturing of the porous structures 04a and 04b there were differences in the particle size fraction and the sintering time. Structure 04b was sintered from bigger particles than 04a, thus explaining the bigger average pore size in 04b. When sintering time is increased the particles will grow closer together leading to
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5 Conclusion
Fig. 9 Production of cytokines a IL-6 and b IL-8 by keratocytes on different porous glass structures and by keratocytes on control TC plastic in response to LPS or no LPS
a more compact structure, thus explaining the lower total porosity of 04b compared to 04a. The sintering parameters were the same for 98a and 04a, but the glass 1-98 used for 98a structure had a lower silica content and thus a lower viscosity during sintering than glass 28-04 in 04a structure leading to lower porosity of 98a than 04a when using the same sintering parameters for both compositions. Cell invasion deeper into the matrix depends on the pore size of the glass structure. The initial cell adhesion results indicate that cells moved deeper into the structures with large pores. The biggest average pore size (135 lm) and trabecular distance (around 300–450 lm) was measured in 04b. In the two other structures with smaller average pore size, the cells were observed closer to the outer surfaces during the initial cell adhesion period. Both the average pore size and chemical composition of the glasses affected cell adhesion. The cell culture time was relatively short because in this paper only acute cell behaviour was investigated. The effect of long term ion release caused by dissolution of the glasses was not studied. It is known that Ca2?,, Na? and K? are important cell signalling ions and concentration changes of these ions in the tissue might change the cell behaviour [29]. Cell adhesion was highest to the 98a bio-glass which also showed greatest bioactivity and levels of Ca2? release. These ionic species are thought to quicken formation of HA surface coating followed by enhanced cell adhesion and deposition of extracellular matrix. This may provide a partial explanation for the greater numbers of keratocytes adhered to the surface of the 98a bio-glass.
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Three bio-glasses of varying bioactivity, 98a, 4a and 4b, were investigated for use as an OOKP skirt substitute. 98a showed the highest rate of calcium ion release, CaO and SiO2 dissolution over 500 h. The pH level rose for up to 100 h and then stabilised. 98a also showed the highest human keratocyte cell adhesion to the surface indicating that a rise in local ionic species may influence rapid cell adhesion. None of the porous bio-active glass structures induced a cytokine driven inflammatory response and the adherent keratocytes showed a typical elongated, spindle shaped morphology suggestive of good adhesive potential. This supports their use as synthetic OOKP skirt in this respect. However, dissolution of the bio-glass over time may destabilise the OOKP indicating that a composite system using a stable backbone structure may be necessary to maintain the PMMA optic following bio-glass chemical degradation. Future in vivo studies will explore the systemic effects of the bio-glasses and the impact of ion dissolution and pH change in the eye. Acknowledgments Academy of Finland (Project no: 114117) is acknowledged for financial support. Jessica Alm is thanked for advice on cell culture on bio-active glasses.
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