examining porous bio-active glass as a potential osteo-odonto-keratoprosthetic skirt material

Examining porous bio-active glass as a potentialosteo-odonto-keratoprosthetic skirt material

Reeta Huhtinen • Susan Sandeman • Susanna Rose • Elsie Fok • Carol Howell •

Linda Froberg • Niko Moritz • Leena Hupa • Andrew Lloyd

Received: 8 October 2012 / Accepted: 28 January 2013 / Published online: 6 February 2013

� 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 pre-

liminary 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. Struc-

tural 2D and 3D analysis of porous structures were per-

formed. 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

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 kera-

toprosthesis anchored within the cornea. The early kera-

toprostheses were fabricated using glass or quartz glass to

provide a convex disc in the central part of the kerato-

prosthesis, 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 kera-

toprothesis 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: gutta-

percha and casein, both natural polymers [1]. Although the

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. Froberg � L. Hupa

Process Chemistry Centre, Abo Akademi University,

Turku, Finland

123

J Mater Sci: Mater Med (2013) 24:1217–1227

DOI 10.1007/s10856-013-4881-x

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 revolu-

tionary 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 keratopros-

thesis 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 (AlphaCor� previously known as Chirila KPro)

[3], polyurethanes (Seoul KPro) [4, 5], polytetrafluoroeth-

ylene (PTFE) Teflon� [6, 7], polyethylene terephthalate

Dacron� [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 sur-

rounding tissue and, especially if used in porous form, it

provides a potential route for infection.

Theoretically, the KPro should integrate with the cor-

neal 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 non-

permeable.

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 col-

lagen [13, 14], fibronectin [15, 16] and laminin [17].

Another way to increase the biocompatibility of the

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 appli-

cations the bio-active glass slowly resorbs and is finally

replaced by newly formed tissue. The influence of glass com-

position on in vitro reactions was recently reported [21]. Bio-

active glasses are interesting material options for OOKPs, as

they bond with living tissue. However, one problem with typ-

ical bio-active glasses is that they resorb with time, thus lim-

iting 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]. Bio-

active glass–ceramic coating on titanium in an ocular envi-

ronment 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 composi-

tions could be of interest in OOKP skirt design.

This study investigated the suitability of porous bio-

active glasses as replacement OOKP skirt materials. The

most suitable glass would be a relatively stable base

1218 J Mater Sci: Mater Med (2013) 24:1217–1227

123

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 kerato-

cytes 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 cul-

tured 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 compo-

sition 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

CaHPO4�2H2O in a Pt crucible at 1360 �C for 3 h. After casting

and annealing, the glasses were remelted to improve homo-

geneity. 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 sim-

ulated aqueous humour zinc precipitates and is thus exclu-

ded. 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)amino-

methane 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 concentra-

tions 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 mea-

sured. 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, dis-

solution 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 6 (5.9) 11 (7.1) 5 (7.6) 22 (23.9) 1 (0.9) 2 (0.9) 53 (53.8)

28-04 5 (4.9) 11.25 (7.2) 6 (9.0) 15 (16.2) 3 (2.6) 0 (0) 59.75 (60.1)

J Mater Sci: Mater Med (2013) 24:1217–1227 1219

123

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 experi-

ment. 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 (Radi-

ometer) using a pHC2401 combination pH electrode. The

solution temperature was measured with a T201 ther-

mometer (Radiometer). Prior to experiments the pH elec-

trode 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

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 high-

speed 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). Auto-

matic 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. There-

after, the software provided direct calculations of mor-

phometric parameters including material volume fraction

(BV/TV), mean trabecular thickness (Tb�Th), trabecular

separation (Tb.Sp), and trabecular number (Tb�N). Non-

metric 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 tra-

becular thicknesses (distribution of distances between the

pores) and the distribution of trabecular separations (dis-

tribution of pore sizes), both are presented in the results

section.

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-) 107 107

Bicarbonate (HCO3-) 20.6 20

Phosphorus (P) 0.62 0.62

Zinc (Zn) 1.59 –

Tris – 50


123

2.5 In vitro assessment of keratocyte adhesion

and cytokine production

2.5.1 Cell culture and characterisation of cytokine

expression profile

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 pas-

saged 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 main-

tenance 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 concen-

tration 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 condi-

tioned 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

content according to manufacturer’s instructions (BD Bio-

sciences). Media control (without cells) was also assessed to

determine any background levels of cytokine.

2.5.2 Cell adhesion and cytokine production in response

to bio-glass exposure

Three of each porous glass structures 98a, 04a and 04b were

placed in a 24 well plate and incubated in 1 ml of supple-

mented 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 lipopoly-

saccharide (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

0

20

40

60

80

100

120

0 100 200 300 400 500

Time [h]

98a

04a

04b

0

10

20

30

40

50

60

0 100 200 300 400 500

Time [h]

98a04a04b

0

2

4

6

8

10

0 100 200 300 400 500

Time [h]

98a

04a

04b

0

2

4

6

8

10

0 100 200 300 400 500

Time [h]

98a

04a

04b

a

b

c

d

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

dissolution of SiO2 from porous glass structures d Cumulative

dissolution of CaO from porous glass structures


123

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.

3 Results

3.1 In vitro dissolution

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

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 car-

bonate equilibrium value during the pH experiment.

3.3 Structural analysis

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.

3.4 Keratocyte adhesion and inflammatory response

The keratocyte cell strain reached 35 cumulative popula-

tion doublings (cpd) prior to senescence of the cultures.

Cells at early and mid-phase cpd showed positive expres-

sion 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 cyto-

kine secretion by the keratocytes was measured over a

period of 48 h in the absence and presence of LPS. Sig-

nificant 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). How-

ever, 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

7,40

7,50

7,60

7,70

7,80

0 100 200 300 400 500

Time [h]

pH

98a

04a

04b

Fig. 2 The pH change caused by the porous glass structures in

simulated aqueous humour as a function of immersion time


123

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 indi-

cating 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 diffi-

cult to mimic the properties of bone in synthetic form.

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

0

2

4

6

8

10

12

14

0,00 50 100 150 200 250 300 350 400 450

Pore size (µm)

98a

04a

04b

Fig. 4 Distribution of pore sizes measured with micro-computed

tomography

0

1

2

3

4

5

6

7

8

9

0,00 100 200 300 400 500 600

Trabecular thickness (µm)

98a

04a

04b

Fig. 5 Distribution of distances between pores measured with micro-

computed tomography


123

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 sol-

uble ionic species release from the chemically degrading

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 inflam-

matory cytokine release.

In vitro dissolution results show that the porous struc-

ture 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 struc-

ture 98a than for the structures 04a and 04b in Fig. 1d.

Structures 04a and 04b have the same chemical composi-

tion, but different porosity. The difference in the dissolu-

tion 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 sur-

rounding 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 struc-

tures 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

0

5

10

15

20

25

30

35

1 2 6 24 30 48

Time in hours

0 ng/mL LPS10 ng/mL LPS

100 ng/mL LPS

1000 ng/mL LPS

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 kerato-

cytes 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 calcein-

am (mean ± SEM, n = 3)


123

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 bio-

active glass structures in such an environment can partly

counterbalance the pH changes, and thus have an antibac-

terial 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 struc-

tures 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

Surface layer

Internal

a b

c

d

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


123

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.

5 Conclusion

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 sys-

temic effects of the bio-glasses and the impact of ion dis-

solution 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.

References

1. Chirila TV, et al. Artificial cornea. Prog Polym Sci. 1998;23:

447–73.

2. Langefeld S, Kompa S, Redbrake C, Brenman K, Kirchhof B,

Schrage NF. Aachen keratoprosthesis as temporary implant for

combined vitreoretinal surgery and keratoplasty: report on 10

clinical applications. Graefe’s Arch Clin Exp Ophthalmol. 2000;

238:722–6.

3. Chirila TV. An overview of the development of artificial corneas

with porous skirts and the use of PHEMA for such an application.

Biomaterials. 2001;22:3311–7.

4. Kim MK, Lee JL, Wee WR, Lee JH. Seoul-type keratoprosthesis.

Arch Ophthalmol. 2002;120:761–6.

5. Bruin P, Meeuwsen EAJ, van Andel MV, Worst JGF, Pennings

AJ. Autoclavable highly cross-linked polyurethane networks in

ophthalmology. Biomaterials. 1993;14:1089–97.

6. Pruitt LA. Fluorocarbon polymers in biomedical engineering.

Encycl Mat Sci Tech. 2008;3216-21.

7. Renard G, Cetinel B, Legeais JM, Savoldelli M, Durand J,

Pouliquen Y. Incorporation of a fluorocarbon polymer implanted

at the posterior surface of the rabbit cornea. J Biomed Mat Res.

1996;31:193–9.

8. Caiazza S, Fanizza C, Mazziotti I, Pintucci S, Tomaino MA.

Light and scanning electron microscopy evaluation of the Dacron

felt as the haptic part of an improved keratoprosthesis. An in vitro

and in vivo study. Clin Mat. 1988;3:33–40.

9. Trinkaus-Randall V, Capecchi J, Sammon L, Gibbons D,

Leibowitz HM, Franzblau C. In vitro evaluation of fibroplasia in

a porous polymer. Invest Ophthalmol Vis Sci. 1990;31:1321–6.

10. Trinkaus-Randall V, Banwatt R, Capecchi J, Leibowitz HM,

Franzblau C. In vivo fibroplasia of a porous polymer in the

cornea. Invest Ophthalmol Vis Sci. 1991;32:3245–51.

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


123

11. Kain HL. The development of the silicone-carbon keratoprothe-

sis. Refract Corneal Surg. 1993;9:209–10.

12. Ciolino JB, Dohlman CH. Biologic keratoprosthesis materials. Int

Ophthalmol Clin. 2009;49(1):1–9.

13. Miyashita H, et al. Collagen-immobilized poly(vinyl alcohol) as

an artificial cornea scaffold that supports a stratified corneal

epithelium. J Biomed Mat Res. 2006;76B:56–63.

14. Xie RZ, Sweeney DF, Beumer GJ, Johnson G, Griesser HJ, Steele

JG. Effects of biologically modified surfaces of synthetic lenti-

cules on corneal epithelialization in vivo. Aust N Z J Ophthalmol.

1997;25(Suppl):46–9.

15. Kobayashi H, Ikada Y. Corneal cell adhesion and proliferation on

hydrogel sheets bound with cell-adhesive proteins. Curr Eye Res.

1991;10:899–908.

16. Pettit DK, Horbett TA, Hoffman AS, Chan KY. Quantitation of

rabbit corneal epithelial cell outgrowth on polymeric substrates

in vitro. Invest Ophthalmol Vis Sci. 1990;31:2269–77.

17. Ohji M, Mandarino L, SundarRaj N, Thoft RA. Corneal epithelial

cell attachment with endogenous laminin and fibronectin. Invest

Ophthalmol Vis Sci. 1993;34:2487–92.

18. Strampelli B. Osteo-odonto-keratoprosthesis. Ann Ottalmol Clin

Ocul. 1963;89:1039.

19. Ricci R, Pecorella I, Ciardi A, Della RC, Di Tondo U, Marchi V.

Strampelli’s osteo-odonto-keratoprosthesis. Clinical and histo-

logical long-term features of three prostheses. British J Oph-

thalmol. 1992;76:232–4.

20. Karlsson KH, Ylanen H, Aro H. Porous bone implants. Ceram

Int. 2000;26:897–900.

21. Zhang D, Vedel E, Hupa L, Aro HT. Predicting physical and

chemical properties of bio-active glasses from chemical

composition. Part III. In vitro reactivity of glasses. Glass Tech

Eur J Glass Sci Tech. 2009;50(1):1–8.

22. Hench LL, Andersson O. ‘‘Bio-active glasses’’ an introduction to

bioceramics. In: Hench LL, Wilson J, editors. Advanced series in

ceramics – vol. 1. Singapore: World Scientific; 1993.

23. Linnola RJ, Happonen RP, Andersson OH, Vedel E, Yli-Urpo

AU, Krause U, Laatikainen L. Titanium and bio-active glass–

ceramic coated titanium as materials for keratoprosthesis. Exp

Eye Res. 1996;63:471–8.

24. Arstila H, Vedel E, Hupa L, Hupa M. Predicting physical and

chemical properties of bio-active glasses from chemical compo-

sition. Part II. Devitrification characteristics. Glass Tech Eur J

Glass Sci Tech. 2008;49(6):251–9.

25. Arstila H, Vedel E, Hupa L, Hupa M. Predicting physical and

chemical properties of bio-active glasses from chemical compo-

sition. Part 2: Devitrification characteristics. Glass Tech Eur J

Glass Sci Tech. 2009;49(6):260–5.

26. Lentner C. Geigy scientific tables, vol. 1. Units of measurements,

Body fluid, Composition of body, and Nutrition. West Galdwell:

Giba-Geigy; 1981.

27. Viitala R, Franklin V, Green D, Liu C, Lloyd A, Tighe B.

Towards a synthetic osteo-odonto-keratoprosthesis. Acta Bio-

mater. 2009;5:438–52.

28. Zhang D, Lepparanta O, Munukka E, Ylanen H, Viljanen MK,

Eerola E, Hupa M, Hupa L. Antibacterial effects and dissolution

behavior of six bio-active glasses. J Biomed Mat Res. doi:

10.1002/jbm.a.32564.

29. Hoppe A, Guldal NS, Boccaccini AR. A review of the biological

response to ionic dissolution products from bio-active glasses and

glass–ceramics. Biomaterials. 2011;32:2757–74.


123

http://dx.doi.org/10.1002/jbm.a.32564

examining porous bio-active glass as a potential osteo-odonto-keratoprosthetic skirt material

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