Hierarchical porous materials for tissue engineering

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Hierarchical porous materials for tissueengineering

BY JULIAN R. JONES*, PETER D. LEE AND LARRY L. HENCH

Department of Materials, Imperial College London, South Kensington campus,London SW7 2AZ, UK

Biological organisms have evolved to produce hierarchical three-dimensional structureswith dimensions ranging from nanometres to metres. Replicating these complex livinghierarchical structures for the purpose of repair or replacement of degenerating tissues isone of the great challenges of chemistry, physics, biology and materials science. Thispaper describes how the use of hierarchical porous materials in tissue engineeringapplications has the potential to shift treatments from tissue replacement to tissueregeneration. The criteria that a porous material must fulfil to be considered ideal forbone tissue engineering applications are listed. Bioactive glass foam scaffolds have thepotential to fulfil all the criteria, as they have a hierarchical porous structure similar tothat of trabecular bone, they can bond to bone and soft tissue and they release silicon andcalcium ions that have been found to up-regulate seven families of genes in osteogeniccells. Their hierarchical structure can be tailored for the required rate of tissue bonding,resorption and delivery of dissolution products. This paper describes how the structureand properties of the scaffolds are being optimized with respect to cell response and thattissue culture techniques must be optimized to enable growth of new bone in vitro.

Keywords: bioactive glass; scaffold; tissue engineering; X-ray microtomography;mineralization

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1. Introduction

The average age of the UK population is increasing as life expectancy increasesand birth rate decreases, with 20% of the population being over the age of 60.The life expectancy of men is now 78.5 years, while that of women is 82.5 years.Unfortunately, 60% of those over 60 are chronically ill. In the UK alone one infour people will die from respiratory disease, while three million people sufferfrom osteoporosis. Osteoporosis reduces bone density and affects everyone tosome degree as they age. The density and strength of bones decrease becausebone resorption occurs faster than new bone is produced. The disease eventuallyleads to collapse or fracture of bones, especially in the hip, wrist, knee and spine.

A common consequence of osteoporosis, arthritis and trauma is the need forskeletal replacements. Current surgical procedures for bone repair aretransplantation or implantation.

Phil. Trans. R. Soc. A (2006) 364, 263–281

doi:10.1098/rsta.2005.1689

Published online 2 December 2005

e contribution of 18 to a Discussion Meeting Issue ‘Engineered foams and porous materials’.

uthor for correspondence ([email protected]).

263 q 2005 The Royal Society

http://rsta.royalsocietypublishing.org/

J. R. Jones and others264


The gold standard in reconstructive surgery is the autograft, which is theharvesting of the patient’s tissue from a donor site and transplantation to thedamaged site. Alternatives are homografts (transplantation from anotherpatient) and xenografts (tissue from a different species, e.g. freeze-dried bovinebone). There are many limitations to using these techniques; autografts have lowavailability and can cause morbidity at the donor site. Homografts carry the riskof disease transmission, bone resorption and rejection, requiring indefiniteadministration of immunosuppressant drugs to the patient. Xenografts are inlarge supply but they have even larger risks of immune rejection, in situdegeneration and disease transmission (Jones & Hench 2001).

The treatment for advanced stage arthritis of the hip joint is a total jointreplacement. However, all orthopaedic implants have a limited lifespan as theylack three of the most critical characteristics of living tissues: (i) the ability toself-repair; (ii) the ability to maintain a blood supply; and (iii) the ability tomodify in response to stimuli such as mechanical load.

As life expectancy increases there is a growing need for an artificial alternativeto an autograft. A paradigm shift from replacement to regeneration of tissuesmay provide a solution (Hench & Polak 2002).

The aim of regenerative medicine is to restore diseased or damaged tissue to itsoriginal state and function, reducing the need for transplants and jointreplacements. There are two strategies to achieve this aim; tissue engineeringand tissue regeneration. Both of these strategies use scaffolds to guide andstimulate growth and differentiation of cells and form tissues (Langer & Vacanti1993; Davies 2000).

2. Bone tissue engineering

An ideal strategy for the tissue engineering of bone is the harvesting of osteogeniccells from the patient, which are then expanded in culture and seeded on ascaffold that acts as guide and stimulus for tissue growth in three dimensions(Ohgushi & Caplan 1999; Takezawa 2003). The osteogenic cells lay down boneextracellular matrix in the shape of the scaffold as woven (immature) bone. Thetissue engineered construct can then be implanted into the patient. Over time,the synthetic scaffold should resorb into the body as non-toxic degradationproducts, allowing the bone to remodel itself into mature bone structure.

(a ) The hierarchical structure of bone

To be able to grow new bone it is important to understand its structure.Figure 1 shows the hierarchical structure of bone. Bone is a natural composite ofcollagen (polymer) and bone mineral (ceramic). Collagen is a triple helix ofprotein chains, a complex structure that has high tensile and flexural strengthand provides a framework for the bone. The helical chains are of the order of10 nm in length and they are arranged into orientated collagen fibres that are ofthe order of 500 nm in length. Bone mineral is a crystalline calcium phosphateceramic (carbonated hydroxyapatite, HCA) that provides the stiffness and highcompressive strength of bone. The mechanism of osteogenesis is complex, butsimplistically the extracellular matrix of mineralizable collagen is laid down byosteoblasts (osteogenic cells), which develop (differentiate) from stem cells

Phil. Trans. R. Soc. A (2006)


size molecules

0.1–1nm nucleic acids

amino acids sugars lipids minerals

macromolecules

10–103nm DNA, RNA proteins polysaccharides bilipidlayers

HCA

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0.1µm nucleus mitochondria RER membrane

4–40µm cells

0.03–103µm extracellular matrix, collagen fibres/ HA crystal composite

1–103µm bone

Figure 1. The hierarchical structure of bone.

265Hierarchical porous materials


and are 20–50 mm in diameter. They secrete type I collagen which thenmineralizes to form an HCA-collagen structure. An osteoblast that becomessurrounded by concentric rings of mineralized tissue is called an osteocyte(figure 1). The two most important types of bone are cortical and cancellousbone. Cortical bone is a dense structure with high mechanical strength and is alsoknown as compact bone. Cancellous or trabecular bone, also called spongy bone,is an internal porous supporting structure of a network of struts (trabeculae)enclosing large voids (macropores). These two types of bone then grow togetherto form a bone, which in the case of a femur is approximately 0.5–1 m long.

Bone is remodelled in response to its local loading environment by the body.Osteoclasts are cells that resorb old bone and bone that is not required (i.e. notunder any load), while osteoblasts lay down new bone. Osteoporosis occurs asosteoblasts become less active but bone is still removed by osteoclasts. The strutsof the trabecular bone are most affected by osteoporosis. An aim of regenerativemedicine is to stimulate the body to reactivate osteogenic cells to re-create thenatural three-dimensional architecture of bone.

When minor damage is done to a bone, it can repair itself by the biochemicalactivity of the osteoblasts, called osteogenesis. However, if the defect exceeds acritical diameter or volume, the bone cannot repair itself. Such defects can resultfrom trauma or from the removal of diseased tissue. Graft implants (transplants)or synthetic bone filler materials are currently used to repair critical size bonedefects. The use of a regenerative scaffold would guide and stimulate bonegrowth. A tissue engineered construct could also be used to fill critical sizeddefects.

3. An ideal scaffold

The general criteria for an ideal scaffold for bone tissue engineering are that it:

(i) acts as template for in vitro and eventually in vivo bone growth in threedimensions;





(ii) resorbs at the same rate as the bone is repaired, with degradationproducts that are non-toxic and that can be easily be excreted by thebody;

(iii) is biocompatible (not toxic) and promotes cell adhesion and activity,stimulating osteogenesis at the genetic level;

(iv) bonds to the host bone without the formation of scar tissue, creating astable interface;

(v) exhibits mechanical properties matching that of the host bone afterin vitro tissue culture;

(vi) is made from a processing technique that can produce irregular shapes tomatch that of the bone defect; and

(vii) has the potential to be commercially producible and sterilizable to therequired international standards for clinical use.

Each of these criteria will be discussed in this paper. In summary, a material isrequired that has properties either very similar to trabecular bone or that can beused to stimulate new bone growth and create a biocomposite that has structureand properties similar to trabecular bone.

To fulfil criterion (i), the scaffold must have an open porous structure to allowcell penetration, tissue ingrowth and eventually vascularization on implantation.There has been much debate regarding the minimum interconnected porediameter to achieve this, and 100 mm is recognized to be the minimuminterconnected pore aperture diameter for an in vivo scaffold as it has beenfound to encourage the vascularization that is required for complete regenerationof bone (Okii et al. 2001). The criteria for an optimized pore network for in vitrobone growth are less clear, especially if the scaffold resorbs in vitro and thestructure changes before implantation. There have been few investigations ofpore orientation, how the pores link to each other to form channels or how fluidflows through the pores to provide a preferential route of cell migration. Thispaper will show how these issues can be addressed.

To achieve criterion (ii), resorbable porous polymeric scaffolds have beendeveloped. Bone cells may initially attach to polymers in vitro, especially ifattachment specific proteins are incorporated on their surface. However,polymers do not bond to bone and do not stimulate cells at the genetic level.Commonly used polymers, such as polyglycolic acid, have the Young modulusmuch lower than bone and many biodegradable polymers degrade rapidly,reducing the strength of the scaffolds before tissue can regenerate. They,therefore, do not fulfil criteria (iii), (iv) and (v) and there is also concern over theacidic degradation products of the biodegradable polymer scaffolds (Holy et al.2003). Bioactive materials have the potential to fulfil criteria (ii) and (iii). Whenimplanted into the body, bioactive materials stimulate a biological response fromthe body such as bonding to tissue (Hench & Wilson 1993). There are two classesof bioactive material; class B bioactive materials bond to hard tissue (bone) andstimulate bone growth along the surface of the bioactive material (osteoconduc-tion). Examples of class B bioactive materials are synthetic hydroxyapatite(HA), tri-calcium phosphate ceramics (b-TCP) and HA-coated porous titaniumoxide (titania). Class A bioactive materials not only bond to bone and areosteoconductive but they are also osteoproductive, i.e. they stimulate the growthof new bone on the material away from the bone/implant interface and can bond





to soft tissues such as gingival (gum) and cartilage. Examples of class A bioactivematerials are bioactive glasses.

4. Bioactive glasses

Bioactive glasses are based on a random network of silica tetrahedra containingSi–O–Si bonds. The network can be modified by the addition of networkmodifiers such as Ca, Na and P, which are bonded to the network via non-bridging oxygen bonds. The mechanism of bone bonding to bioactive glasses isdue to the formation of a carbonate substituted hydroxycarbonate apatite layer(HCA) on the surface of the materials after immersion in body fluid. This layer issimilar to the apatite layer in bone and, therefore, a strong bond can form (Hench1991). Bioactive glasses are osteoproductive, which means they stimulate newbone growth on their surface, even away from the glass/bone interface(osteoproduction) and can be resorbable (Hench & Polak 2002).

Importantly for criterion (iii), the dissolution products of bioactive glasses(soluble silicon and calcium) have been found to up-regulate seven families ofgenes in osteoblasts (Xynos et al. 2000a–c) and to have an effect on the cell cyclewhereby more cells express biochemical markers for new bone formation and cellsthat are not capable of forming new bone are eliminated by apoptosis(programmed cell death; Hench et al. 2000). Synthetic HA does not releasesuch dissolution products. However, chemical substitution of silicon for calciumin synthetic HA shows improved bone ingrowth in vivo over phase pure HAgranules (Patel et al. 2002).

There are two types of bioactive glasses; melt-derived and sol-gel derived.A certain composition of melt-derived bioactive glass (46.1 mol% SiO2,24.4 mol% Na2O, 26.9 mol% CaO and 2.6 mol% P2O5), called Bioglass is usedin the clinic as a treatment for periodontal disease (Perioglas) and as a bonefilling material (Novabone; Hench 1991; Fetner et al. 1994). Bioglass implantshave also been used to replace damaged middle ear bones, restoring the hearingto thousands of patients and as tooth root replacements (Wilson et al. 1995).

Sol-gel derived bioactive glasses are synthesized by the hydrolysis of alkoxideprecursors to form a sol, which is a colloidal silica solution. The sol thenundergoes polycondensation to form a silica network (gel). The gel is then heattreated to form a glass (Brinker & Scherer 1990; Hench & West 1990). Sol-gelderived bioactive glasses tend to be more bioactive and resorb quicker than melt-derived glasses of similar compositions. This is because sol-gel glasses have ananometre scale textural porosity that is inherent to the sol-gel process, whichincreases the specific surface area by two orders of magnitude compared to amelt-derived glass of a similar composition (Sepulveda et al. 2002a). The texturalporosity not only increases the surface area for cation exchange and networkdissolution by two orders of magnitude, but it also exposes many silanol groupsto the solution, which act as nucleation sites for HCA layer formation.

(a ) Bioactive glass scaffolds

Pores have been introduced into melt-derived bioactive glasses but the poreswere few in number and were in the form of orientated channels of irregulardiameter running through the glass so interconnectivity was poor (Yuan 2001).



sol-preparation from a mixture of alkoxides(distilled water, HNO3, TEOS, Ca(NO3)2, depending

on glass composition required)

ageing at 60˚Cdrying at 130˚C

thermal stabilization at 600–800˚C

+ gelling agent + surfactantfoaming by vigorous agitation

casting into sealable moulds

further sintering at 700, 800, 1000˚C

Figure 2. A flow diagram of the process for producing bioactive glass foam scaffolds.



By foaming sol-gel derived bioactive glasses, our team at Imperial CollegeLondon, produced scaffolds with a hierarchical pore structure similar totrabecular bone (Sepulveda et al. 2002b).

Figure 2 shows a flow chart of the sol-gel foaming process. In the first step a solis synthesized from a silica-based alkoxide precursor, such as tetraethyloxysilane(TEOS).

After hydrolysis is completed, the sol is foamed by vigorous agitation in air.The viscosity increases rapidly due to the addition of a gelling agent(hydrofluoric acid, HF) and a surfactant is added, which lowers the surfacetension and stabilizes the air bubbles on initial foaming (Rosen 1989). Thebubbles are permanently stabilized by the gelation reaction (polycondensation).Tertiary (SiO2, CaO, P2O5), binary (SiO2, CaO) and unary systems (SiO2) canall be successfully foamed as scaffolds (Sepulveda et al. 2002b).

(b ) Hierarchical structure of bioactive glass scaffolds

Figure 3 shows a scanning electron micrograph (SEM) of a bioactive glassfoam of the 70S30C (70 mol% SiO2, 30 mol% CaO) composition. Figure 3 showsthat the foam is comprised of large macropores with diameters in the region of200–600 mm that are highly interconnected (dark areas). Many of the apertureshave diameters in excess of the 100 mm required for tissue engineeringapplications. As the foam is made from sol-gel derived bioactive glass, thethree-dimensional interconnected solid network has a textural porosity withdiameters in the range 2–20 nm, termed mesoporosity.

Figure 4 shows an X-ray micro-computed tomography (XMT) image of asimilar scaffold, to that shown in figure 3, obtained using a commercial XMT unit(Phoenix X-ray Systems and Services GmbH). The XMT unit is based on thesame principles as a CAT scan (computed axial tomography) where series of two-dimensional transmission X-ray images are reconstructed to form a three-dimensional image. The key difference is that geometric enlargement is used to



Figure 4. XMT image of a bioactive glass foam scaffold.

Figure 3. SEM image of a typical bioactive glass foam scaffold.



magnify the image by placing the object close to a micron sized spot source,producing a magnified image which is projected onto a solid-state detector alarge distance from the object (relative to the source–object distance). Like aCAT scan, it provides quantitative data on the integrated density and atomicnumber of the matter in each voxel (volume pixel). Reconstructed imagesconsisting of 512!512!512 voxels, each of 4.7 mm on a side, were collected fromeach sample set and cropped digitally to remove the edge artefacts. Figure 4shows that the macropore network is very highly interconnected and is verysimilar to the XMT image of trabecular bone shown in Stock (1999). The pore



Figure 5. XMT of an interconnected single pore of bioactive glass foam scaffold.



shape and size appear to be very homogeneous. This is because the pores formin a liquid that is well mixed with an evenly dispersed surfactant content. Infigure 5, an isolated pore that is representative of the entire sample, which isconnected to other pores on 4–6 sides. Connectivity occurs because the sphericalair bubbles are all in contact with each other immediately prior to gelation,separated only by a thin film of silica-based sol that is stabilized by thesurfactant. Upon gelation and subsequent thermal processes the thin film drains,shrinkage occurs and the surfactant is combusted, leaving the apertures.

(c ) Tailoring of the structure

We have found that variables in each stage of the foaming process (figure 2)have an effect on the pore structure. Such variables include the sol (glass)composition and surfactant concentration (Jones & Hench 2003), gelling agentconcentration, the temperature at which the process is carried out and whetheradditional water is added with the surfactant to improve its efficiency (Joneset al. 2004a). For specific applications it may be necessary to select a particularpore diameter and interconnected pore size. For tissue engineering applicationsthe macropore diameter has little importance, but the modal interconnected porediameter should be greater than 100 mm. Changing the surfactant concentrationwhile keeping all other variables constant is the most efficient method to controlthe aperture diameter (Jones & Hench 2003).

Figure 6 shows the dissolution profiles of silicon, calcium and phosphate ionsfrom powders, foams and monolithic discs of bioactive glasses of the 58S(60 mol% SiO2, 36 mol% CaO, 4 mol% P2O5) composition. The same mass of



foam monoliths powder

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each glass was immersed in simulated body fluid (SBF) and the ions releasedwere quantified by inductive coupled plasma analysis (ICP). Figure 6 shows thatthe release of the gene activating ions Si and Ca was highest for powders andlowest for monoliths. This is because dissolution is more rapid as the ratio ofsurface area to solution volume ratio increases. The amount of phosphate insolution decreases because of formation of a calcium phosphate (HCA) layer onthe surface of the glass. The rate of formation of this layer increased as thedissolution rate increased. The morphology of a potential scaffold will, therefore,have a large effect on rate of delivery and concentration of gene stimulating ionsand therefore rate of bone bonding and regeneration of bone.

None of the processing variables listed above have a large effect on the texturalmesoporosity. However, this can be controlled by changing the final sinteringtemperature of the scaffolds (Jones et al. 2004b).When calciumnitrate is used in thesol-gel process to introduce calcium into the glass composition, the residual nitratesmust be removed to chemically stabilize the glass and make it biocompatible(non-toxic to cells). Nitrates are burnt off at approximately 550 8C, therefore 600 8Cis the minimum sintering temperature for these glasses. Figure 7 shows texturalpore size distributions obtained from foams sintered at 600, 700, 800 and 1000 8C for2 h, using the Barrett Joyney Halenda (BJH) method on data obtained fromnitrogen adsorption analysis (Barrett et al. 1951). The vertical axis is a derivative ofthe volume of nitrogen desorbed from the foam relative to the pore diameter.

Figure 7 shows that scaffolds sintered at 600 and 700 8C exhibit narrow poresize distributions with narrow pore diameters and a modal pore diameter ofapproximately 17 nm. The shape of the pores is difficult to ascertain, but theisotherms (plots of volume of nitrogen adsorbed and desorbed as a function ofpressure at constant temperature, 77 K) were type IV isotherms according toIUPAC classification (Sing), with type II hysteresis loops, which implies ink-bottle shaped pores in the mesopore range (Saravanapavan & Hench 2003).However, this shape of isotherm would also be obtained for cylindrical pores thatcontain bulbous portions. Figure 6 shows that as the sintering temperatureincreased from 700 to 800 8C, the modal pore diameter dropped to approximately12 nm and the textural porosity appeared to have been removed after sintering at1000 8C. The decrease in textural porosity coincides with an increase incompressive strength. Foams sintered at 600 8C have a compressive strength ofapproximately 0.25 MPa (Instron parallel plate, diameter to height ratio offoam discs was 3 : 1) while similar foams sintered at 800 and 1000 8C have acompressive strength of approximately 2.5 MPa (Jones et al. 2004b), similar tothat of trabecular bone (Hench & Wilson 1993). Figure 8 shows XMT images ofsections of scaffolds sintered at 700, 800 and 900 8C. The images show that themacropore diameter also decreased with sintering, but it is important to quantifyhow the interconnected pore diameter was affected. However, obtainingquantitative data from XMT images has required developing new interpretivesoftware (Atwood et al. 2004).

5. Quantitative three-dimensional image analysis

Although the XMT images provide quantitative data of the integrated densityand composition of the scaffolds at evenly spaced points in three dimensions,



Figure 8. Images obtained from XMT of sections of scaffolds sintered at 700, 800 and 900 8C for 2 h.



converting these pictures from greyscale images to quantified descriptors of thestructures requires the development and use of appropriate mathematicalmorphological operators. As an example, a single two-dimensional slice from athree-dimensional scan of a 70S30C scaffold is shown in figure 9a. The apertureswhich provide the interconnectivity between the pores (black) at areas in thescaffold walls (whitish) are visible; however, quantifying the size of theseapertures requires classification of the image into individual pores. This is simpleif the pores are closed, but for this open cell porosity a new algorithm wasdeveloped (as described in detail in Atwood et al. 2004):

(i) threshold the image, classifying each voxel as either scaffold or empty space;(ii) apply a dilation algorithm to grow from the scaffold walls into the centre of

the pores, noting the number of steps it has taken to grow to each voxel.Centroids of each pore will fill last;

(iii) using the centroids, together with the number of steps grown as a distancemap, a three-dimensional watershed algorithm (Mangan &Whitaker 1999)was applied to divide the image into individual pores. (Watershedalgorithms find the set of points in a function, considered as a height map,that divide regions in whichwater flows to the same final point; analogous tothe watersheds of a river basin in geography.);

(iv) voxelswith neighbours on the same twoporeswere then grouped anddefinedas apertures; and

(v) the individual pores and apertures objects were then quantified to determinetheir volume/area and maximum diameter.

Using this algorithm, the scaffolds shown in figure 8 were quantified and thepore and aperture size distributions are plotted in figure 10. The pore sizedecreases with increasing firing temperature; however, the median interporeaperture diameter does not change significantly.

Therefore, after sintering at 800 8C for 2 h, the scaffold has a modalinterconnected pore diameter in excess of 100 mm and a maximum compressivestrength of 2.4 MPa.

6. Permeability

During in vitro cell culture, the flow of culture medium containing cells duringcell seeding is critical to developing an evenly populated scaffold. The flow in



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Figure 9. Images of two-dimensional slices of a scaffold (a) initial two-dimensional slice, (b) thresholdimage and (c) the same slice after watershed algorithm has been applied to separate pores.



porous media has been shown to be described at a macroscopic level by thegeneralized tensor form of Darcy’s law, allowing the bulk velocity to be related tothe change in pressure using what is called the permeability tensor, �K . Therefore,�K provides a quantitative descriptor of the ease at which seeded cells will



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Figure 11. (a) Reconstruction of scaffold with calculated streak-lines showing where the flow willoccur. (b) Permeability tensor calculated from the reconstruction geometry, illustrating thatminimum of a 1 mm cube of material is required to obtain representative flow.

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penetrate the scaffold, as well as the ease of getting nutrient fluid into the scaffoldduring tissue growth. Measuring the permeability can be difficult since flow willoccur around the edge of the sample as well as through it, and preference flowchannels may form as material dissolves under the flow rates required formeasurement. However, an alternative method for measuring the permeabilityexists; using the three-dimensional geometry of the scaffold obtained via XMT ina microscale flow simulation. The flow within porous medium obeys Stokesequations at the local scale (Sahimi 1995), hence the permeability can becalculated using the geometry and by numerically solving Stokes equations(Papathanasiou & Lee 1997). For this study the permeability was calculatedusing the code developed by Prof. Bernard (CNRS Bordeaux) to study waterflow in reservoir rocks (Anguy et al. 1995).

The resulting flow predictions on a local scale within the complex three-dimensional structure are shown in figure 11a. The plotted streaklines illustratethat the flow is dominated by the size of the apertures. The calculatedmacroscopic components of the permeability tensor are plotted in figure 11b as a





function of the size of the representative volume element (RVE) used in thesimulation. For a RVE of the size of the pores (350 mm), the x, y and zcomponents of �K are very different. However, once the RVE is larger then threepores across (1 mm), the different components converge. This illustrates that the70S30C composition scaffolds fired at 800 8C have an isotropic structure and thatonly a volume of approximately 1 mm3 need be simulated to determine the flowcharacteristics to design the seeding and nutrient flow techniques.

7. Mechanical properties

Measuring the mechanical properties of the scaffolds as a function of the scaffoldpore structure and soaking time is laborious, and it cannot provide a tool forextrapolation to design an optimal scaffold structure. An alternative is to extractthe internal structure of the scaffolds as surfaces from the XMT images and meshthe volumes created by these surfaces. This was performed using scaffolds of the70S30C composition fired 800 8C, with the resulting mesh shown in figure 12a.Using bulk properties for the scaffold material from the literature (Amaral et al.2002), the resulting structure was compressed by displacing the nodes along thetop face downwards whilst fixing the nodes on the bottom face. The simulatedload versus displacement graph was then converted into an effective stiffness forthe porous structure, predicting an effective modulus of 3.8 GPa, in reasonableagreement with a modulus of 3.2 GPa measured on a similar sample. Althoughonly a single scaffold was analysed, it illustrates the viability of implementingsuch a methodology which can be used to design scaffold materials (the bulkstiffness can be tuned by either altering the composition or the mesoporosity) andstructures which exhibit mechanical properties matching that of the host bone,at least after in vitro tissue culture.

8. Molecular level characterization

A collaboration with the teams of Professors Mark Smith (University ofWarwick) and Bob Newport (University of Kent) has allowed the characteriz-ation of the glass network structure at the atomic level using magic anglespinning (MAS), nuclear magnetic resonance (NMR) and X-ray diffraction.Solid-state NMR is an element specific probe technique with high sensitivity tothe local structural environment around the nucleus under investigation(MacKenzie & Smith 2002). 43Ca MAS NMR spectra have been obtained forunstabilized samples of sol-gel derived calcium silicates (70S30C) heated at 120and 350 8C, with the spectra suggesting that at this stage the calcium remainedin an environment similar to the initial calcium nitrate (Lin et al. 2004). As thetemperature was increased the samples became more disordered and no calciumsignal was observed.

Extended X-ray absorption fine structure spectroscopy (EXAFS) and X-rayabsorption near edge structure (XANES), X-ray fluorescence spectroscopy (XFS)and X-ray powder diffraction (XRD) were also used to study the local calciumenvironment in sol-gel-derived bioactive calcium silicate glasses (Skipper et al.2004). The calcium oxygen environment was found to be six-coordinate across arange of binary compositions. The formation of the HCA layer on the 70S30C



Figure 13. SEM image of a mineralized bone nodule on the surface of a pore in a bioactive glassfoam (58S composition) after two weeks in culture with no supplementary factors. Courtesy of DrJulie Gough.



composition in SBF was also investigated. Both the EXAFS and XANES showeda gradual increase in coordination number and Ca–O bond distance as immersiontime in SBF increased. XFS showed that calcium was quickly lost from the glasson exposure to SBF. XRD showed that the formation of the crystalline HCAlayer was preceded by formation of a non-crystalline calcium phosphate phaseafter 1 h of immersion in SBF.

These techniques provide the basis for relating the mechanisms of dissolutionand bioactivity to the molecular structure of the porous bioactive glass scaffolds.The results show that it is important to be able to monitor the molecularstructure of the scaffolds as well as the hierarchical pore structure in order tooptimize the scaffolds from the molecular to macro scale.

9. Cell response

Primary human osteoblasts, harvested from the tops of femurs removed duringtotal hip replacements, have been cultured on bioactive glass foams of both the58S and 70S30C compositions. Gough et al. (2004) seeded the cells (passage twoor three) on to foams of the 58S composition at seeding density of80 000 cells cmK3. Prior to culture the foams were mounted in agar, sterilizedwith UV light and soaked in culture medium for 3 days. During cell culture themedia was changed every 2 days. The cells attached, proliferated and formedsecreted bone extracellular matrix (mainly collagen type I), which mineralizedafter 10 days of culture. Mineralization is the development of HCA, bonemineral. Figure 13 shows an SEM image of primary human osteoblasts culturedon a 58S foam for 10 days. The seeded scaffolds were fixed, dehydrated and goldcoated and observed in an SEM at 15 kV. Figure 13 shows an SEM image of amineralized bone nodule inside a macropore. A bone nodule is a group of cellsthat have laid down some extracellular bone matrix. Mineralized bone nodulescan form on many biocompatible materials in vitro, if supplementary growthfactors, such as dexamethasone, are added to the medium. When cultured onthe bioactive glass scaffold, the nodules mineralized without the addition of



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Figure 14. Graph of relative metabolic activity as a function of mean pore diameter for A549human carcinoma cells after 48 h culture on silica discs.



mineralization supplements to the culture media, which indicates great potentialof the bioactive glass foam for use as an osseous tissue scaffold. The release ofcalcium and silicon ions from the glass are thought to stimulate the rapidmineralization.

The role of phosphate in the glass scaffold composition has much less effect.When cells were cultured on 70S30C scaffolds, mineralized bone nodules wereobserved after two weeks of culture without supplements. This result impliesthat combinations of silicon and calcium ions released from the scaffold stimulatethe cells.

The effect of the surface texture of porous silica (100S) on the proliferation oflung cells was investigated by culturing cells from the A549 (human lungcarcinoma) cell line on sol-gel derived monoliths with different mean porediameters in the range 25–740 A (2.5–74 nm). A cell seeding density of20 000 cells cmK2 was used. After 48 h in culture the cells were fixed withparaformaldehyde. A primary antibody for vinculin conjugated to thefluorochrome Fluorescein was used for immunoflorescence staining. Positivelystained cells were counted from five randomly selected fields of view at 10!10magnification using a florescence microscope. Figure 14 shows a graph of relativemetabolic activity as a function of mean pore diameter of silica discs, where therelative metabolic activity is the percentage number of cells positively stained forvinculin within a field of view in the florescence microscope compared to thenumber on the control. The control is dense fused silica glass (pore diameter ofzero). Figure 14 shows that the proliferation of the lung cells increased as meanpore diameter of the substrate increased, up to maximum at a mean porediameter of 75 A (7.5 nm). It is not clear why 7.5 nm is the optimum porediameter for lung cell attachment and growth, but the results show thatoptimizing the textural porosity of sol-gel derived bioactive glass scaffolds isimportant for optimal cell response. The optimal mesoporosity is likely to becorrelated with the biochemistry of the cell membrane cytokines that areinvolved in cell attachment.





10. A hybrid scaffold for in situ bone regeneration

An optimized bioactive glass foam has high potential to be used as a scaffold togrow a bone/scaffold biocomposite with tissue engineering techniques; however,it would have too low a strength in tension to be used as a scaffold for in situ boneregeneration. In such an application a scaffold (with or without cells seeded on it)would be implanted directly into a defect site without culturing new tissue on thescaffold first. In this case the mechanical properties of the scaffold are highlyimportant. First, the majority of defect sites in bone are load-bearing sites.Secondly, in this case the Young modulus of the scaffold should match that of thehost bone to prevent stress shielding. It is therefore necessary to mimic thehierarchical structure of trabecular bone as a whole, rather than just the bonemineral, i.e. a polymer must be introduced in the ceramic scaffold structure tomimic the collagen/mineral composite of bone, but the interconnected porenetwork and bioactivity of the scaffold must be maintained.

The foaming process was modified, to create bioactive glass/polymer hybridscaffolds, by reacting poly(vinyl alcohol) (Acros Organics, USA, averagemolecular weight of 16 000) in acidic solution with tetraethylorthosilicate. Theinorganic phase was also modified by incorporating a calcium compound.Hydrated calcium chloride was used as precursor. The polymer/sol was thenprocessed in the same way as in figure 2, except that the gelled foam hybrids wereaged at 40 8C and vacuum dried at 40 8C. Mechanical behaviour of the hybridmaterials produced was determined by compression test (parallel plate method)using dynamic mechanical analysis (DMA) equipment. From SEM images, thepore network of the hybrid foams was very similar to that of the glass foams.

Hybrid foams of the glass composition 70S30C containing 20 wt% PVA werecompression tested using DMA parallel plate compression. Sample dimensionswere 5!5!5 mm at a loading rate of 0.5 N minK1. The hybrid foams exhibitedhigher compressive strength and higher deformation to failure than the glassfoams with similar porosity, therefore hydrid glass/polymer foams have thepotential to have mechanical properties suitable for implantation into loadbearing defect sites.

11. Conclusions

A bioactive glass scaffold of the 70S30C composition has the potential to serve asa scaffold for bone tissue engineering applications. X-ray microcomputertomography can be used in conjunction with three-dimensional image analysisto quantify the macropore network and to non-destructively predict fluid flowwithin the scaffold and its mechanical properties. A compressive strength of2.4 MPa can be attained by sintering the macroporous scaffold at 800 8C for 2 hand the modal interconnected pore diameter remains to be in excess of 100 mm.The techniques of NMR, EXAFS, XANES, XFS and XRD provide the basis forrelating the mechanisms of dissolution and bioactivity to the molecular structureof the glass scaffolds. Optimization of the scaffolds from the molecular to themacro scale with respect to cell response is vital if an ideal scaffold is to bedeveloped. When primary human osteoblasts are cultured on the bioactive glassfoam scaffolds mineralized bone nodules form within 10 days of culture without





the addition of supplementary growth factors. Bioactive glass/polymer hybridscaffolds have the potential to fulfil all the criteria for an ideal scaffold for bothbone tissue engineering and in situ bone regeneration.

The authors thank US Defence Advanced Research Projects (Contract no. N66001-C-8041),

EPSRC, MRC, Lloyds Tercetenary Foundation and the Royal Academy of Engineering for

financial support. The authors would also like to thank Dr Robert Atwood (3D image analysis), Dr

Dominique Bernard (permeability model), Dr Daan Maijer (mechanical property model), Olga

Tsigkou, Papy Embanga and Dr Molly Stevens (all cell biology) for their valuable assistance.

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Hierarchical porous materials for tissue engineering

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