WO2014124496A1 - Matériau biocompatible et utilisations de celui-ci - Google Patents

Matériau biocompatible et utilisations de celui-ci Download PDF

Info

Publication number
WO2014124496A1
WO2014124496A1 PCT/AU2014/000126 AU2014000126W WO2014124496A1 WO 2014124496 A1 WO2014124496 A1 WO 2014124496A1 AU 2014000126 W AU2014000126 W AU 2014000126W WO 2014124496 A1 WO2014124496 A1 WO 2014124496A1
Authority
WO
WIPO (PCT)
Prior art keywords
bioglass
composite material
polymer
sol
hybrid
Prior art date
Application number
PCT/AU2014/000126
Other languages
English (en)
Inventor
Fariba Dehghani
Roya RAVARIAN
Wojciech Chrzanowski
Original Assignee
The University Of Sydney
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Priority claimed from AU2013900475A external-priority patent/AU2013900475A0/en
Application filed by The University Of Sydney filed Critical The University Of Sydney
Publication of WO2014124496A1 publication Critical patent/WO2014124496A1/fr

Links

Classifications

    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L31/00—Materials for other surgical articles, e.g. stents, stent-grafts, shunts, surgical drapes, guide wires, materials for adhesion prevention, occluding devices, surgical gloves, tissue fixation devices
    • A61L31/12—Composite materials, i.e. containing one material dispersed in a matrix of the same or different material
    • A61L31/125—Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix
    • A61L31/127—Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix containing fillers of phosphorus-containing inorganic materials
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
    • A61L27/40—Composite materials, i.e. containing one material dispersed in a matrix of the same or different material
    • A61L27/44—Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix
    • A61L27/46—Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix with phosphorus-containing inorganic fillers
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
    • A61L27/50—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
    • A61L27/56—Porous materials, e.g. foams or sponges
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L31/00—Materials for other surgical articles, e.g. stents, stent-grafts, shunts, surgical drapes, guide wires, materials for adhesion prevention, occluding devices, surgical gloves, tissue fixation devices
    • A61L31/14—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
    • A61L31/146—Porous materials, e.g. foams or sponges
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2400/00—Materials characterised by their function or physical properties
    • A61L2400/06—Flowable or injectable implant compositions
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2430/00—Materials or treatment for tissue regeneration
    • A61L2430/02—Materials or treatment for tissue regeneration for reconstruction of bones; weight-bearing implants
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2430/00—Materials or treatment for tissue regeneration
    • A61L2430/06—Materials or treatment for tissue regeneration for cartilage reconstruction, e.g. meniscus
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2430/00—Materials or treatment for tissue regeneration
    • A61L2430/12—Materials or treatment for tissue regeneration for dental implants or prostheses
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2430/00—Materials or treatment for tissue regeneration
    • A61L2430/24—Materials or treatment for tissue regeneration for joint reconstruction
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L2430/00—Materials or treatment for tissue regeneration
    • A61L2430/38—Materials or treatment for tissue regeneration for reconstruction of the spine, vertebrae or intervertebral discs

Definitions

  • the present invention relates to the synthesis and application of novel organic-inorganic composite molecularly-coupled "hybrid" materials.
  • the hybrids of the invention have applications in wide-ranging technologies such as optics, electronics, mechanics, membranes, protective coatings, catalysis, sensors, and in particular as biomaterials.
  • These hybrids embody the advantageous properties of both organic and inorganic frameworks such that the chemical coupling between these components yields properties that are not readily achievable in conventional composite materials such as thermoplastics and bioglasses, and mere physical admixtures thereof.
  • the inventive materials of the invention can be fabricated from a wide range of different types of polymers including natural, synthetic, degradable and non- degradable. Different types of structures can be produced such as monoliths, fillers, coatings and injectable pastes/solutions which are hardenable in situ.
  • the bioactivity of polymeric implants can be improved by the impregnation of these bioactive materials into such implants, or providing a surface coating thereon to improve biocompatibility and mechanical properties.
  • the hybrid materials of the invention can also formulated as drug delivery devices capable of delivering hydrophobic or hydrophilic pharmaceutical compounds.
  • Bioceramics such as "Bioglass®” are favourable materials for bone grafting applications due to a relatively high biocompatibility and bonding affinity to the host tissue of the mammalian body via the formation of a biologically active hydroxyl carbonate apatite (HCA) layer on their surfaces.
  • HCA biologically active hydroxyl carbonate apatite
  • This HCA phase is chemically and structurally equivalent to the mineral phase in bone, thereby facilitating interfacial bonding.
  • a thin layer of apatite forms on the glass-tissue interface, facilitating strong bond to the bone.
  • formulations can also facilitate growth of osteoblasts through the material. It will be appreciated that reference to "bioglass” throughout the present application is intended to be non-limiting, and any biocompatible glass source is within the scope of the present invention.
  • Bioglass® is a commercially-available family of bioactive glasses, comprising Si0 2 , Na 2 0, CaO and P 2 0 5 in specific proportions. Bioglasses differ from traditional soda-lime glasses due to the relatively low amount of silica (i.e., less than 60 mol.%), relatively high amount of sodium and calcium - and relatively high calcium : phosphorus ratio. The ratio of calcium to phosphorus promotes the formation of apatite crystals; calcium and silica ions can act as crystallisation nuclei.
  • Bioglasses have many different formulations. Some bind to soft tissues and bone (e.g., 45S5), some to bone only (e.g., 5S4.3 or Ceravital), some do not form a bond at all and after implantation become encapsulated with non- adhering fibrous tissue, and others are completely resorbed within few weeks.
  • bioglasses having ⁇ 35 mol.% Si0 2 are said to be non glass-forming.
  • Those having >50 mol.% Si0 2 , ⁇ 10 mol.%) CaO, ⁇ 35 mol.%> Na 2 0 are said to be bioactive, and undergo resorption within 10-30 days.
  • bioglasses having >65 mol.%> Si0 2 are non-bioactive, nearly inert and become encapsulated with fibrous tissue.
  • one of the major obstacles to the widespread adoption of bioglass in bone repair therapies has been its relatively poor mechanical properties, i.e. high Young's modulus and low fracture resistance, which make these materials brittle.
  • the bending strength of bioglass is typically within the range of 40-60 MPa, which is insufficient for load-bearing applications.
  • PMMA poly(methyl methacrylate)
  • PMMA bone cement may be used to affix implants and to remodel lost bone.
  • PMMA-based bone cements were first introduced in the 1960s (see, e.g., Charnley, J. Bone Joint Surg. Br., 1960, 42-B, 28-30) and in the interim, progress has been made toward improving the properties of such materials for the fixation of implants and filling bone voids after trauma or the removal of tumours.
  • MM A methyl methacrylate
  • initiators and activators are placed in the body and fixation occurs by in situ polymerisation.
  • This approach is fraught with several issues that include damage to/necrosis of the surrounding tissues as a result of temperature increase (c. 70°C) from the resultant exothermic reaction; release of toxic compounds such as residual MMA monomer, initiators and activators; inertness and lack of bioactivity, which leads to the thickening of intervening fibrous tissue layer; inhibition of cell function and growth/differentiation; increased inflammatory response; and cell death/necrosis, which can lead to the loosening of prosthesis/implants.
  • PMMA cement Another major consideration when using PMMA cement is the effect of stress shielding. Since PMMA has a Young's modulus greater than that of natural bone, the stresses are loaded into the cement and thus, the bone no longer receives the mechanical signals to continue bone remodelling. This, in turn, can result in bone resorption.
  • Bioglass and PMMA have been physically combined in the past, and several of the abovementioned problems addressed.
  • embedding PMMA into bioglass structure improves certain mechanical properties and reduces brittleness, which in turn provides a material useful in biomedical applications such as bone implants.
  • bioglass has been added in the form of powdered filler to PMMA-based biomaterials to enhance its biocompatibility and bioactivity.
  • the amount and particle size of the bioglass powder has a significant effect on the bioactivity of the resultant composite. Decreasing the particle size of the bioglass increases the available surface area of the bioglass, which results in increased exposure of the bioactive compounds to the surroundings.
  • a physical mixture of PMMA and bioglass in solid form results in a non-homogenous distribution, and a lack of adhesion between the phases which consequently enhances the risk of implant failure.
  • a non-homogenous mixture of PMMA and bioglass results in inconsistent physical properties such as degradation rate and mechanical strength, which is an issue inherent in many commercially-available materials.
  • addition of bioceramic particles such as bioglass to injectable formations of PMMA cements results in needle blockage, which makes the injection difficult and painful for the patient. Also, it results in inhomogeneous filling of the defected site.
  • the present invention provides chemical bonding between organic and inorganic compounds in organic-inorganic molecularly-coupled "hybrid" materials, which exhibit synergistically-enhanced properties.
  • Chemical coupling e.g. covalent bonding
  • the bioactive glass-polymer hybrid materials of the invention are composites consisting of two constituents which are bonded at the nanometer or molecular level.
  • the present invention provides nanoscale interpenetrating networks of the bioglass and polymer which have covalent coupling between them. In some embodiments, this involves careful control of the chemistry of the sol-gel process.
  • the materials of the present invention have potential in wide-ranging technologies such as optics, electronics, mechanics, membranes, protective coatings, catalysis, sensors and biomaterials.
  • the polymer is PMMA, however in alternative embodiments, the polymer may be poly(vinyl alcohol), chitosan, or other biologically compatible polymers, and combinations thereof, to produce bioactive glass hybrid scaffolds for biomedical applications.
  • the physico-chemical properties of the materials of the invention can be "tuned” by adjusting the relative ratios of the individual components, and thereby provide an opportunity to exploit the various properties of both organic and inorganic frameworks, as well as manipulate their functional characteristics by varying parameters such as composition, procedure, coupling and the degree of dispersion of each component within the other.
  • the uniform intermixing of the organic and inorganic phases to provide an interpenetrating network yields unique properties that are as yet unachievable in conventional composites, or even nanocomposites.
  • the organic-inorganic hybrids of the invention additionally address the issue of phase separation, which is a common problem in the prior art, by creating interaction at a molecular level.
  • the organic and inorganic components are mixed in domain sizes of nanometers, and the molecular interactions between the phases lead to the formation of two distinct classes of hybrid.
  • Class I hybrids include a mixture in which there is a weak interaction between the organic and inorganic phases, such as van der Waals forces or hydrogen bonding, whereas “class II” is characterised by hybrids having covalent or iono- covalent bonds, which impacts upon the final material properties.
  • the bonding between phases in the organic-inorganic hybrids of the invention can be class I and/or class II.
  • the present invention relates to organic-inorganic hybrids, wherein the inorganic component is bioglass, and the organic component comprises a polymer, pre-polymer, or an oligomer, or combinations thereof.
  • a hybrid material is formed when there is chemical coupling of the organic and inorganic components, and in one preferred embodiment a coupling agent is used to covalently couple the organic and inorganic phases together and provide an interpenetrating network of each component.
  • the coupling agent is an organic moiety attached to the organic component which allows chemical bonding of the organic component to the inorganic network.
  • an organosilane group can be provided on the organic component which acts as a "network modifier", since in the final structure the inorganic network is only modified by the organic group.
  • a preferred organosilane coupling agents is 3-(trimethoxysilyl)propyl methacrylate ("MPMA”), which comprises both organic and inorganic moieties, thereby providing means to covalently bond the respective organic and inorganic components.
  • MPMA 3-(trimethoxysilyl)propyl methacrylate
  • the organic polymer component contains a certain percentage of MPMA.
  • the polymer can be grafted with trimethoxysilyl functionality. Whilst a preferred network modifier or coupling agent is MPMA, it will be appreciated that other coupling agents can be utilised, as discussed further below.
  • the polymer is functionalised with moieties which provide covalent coupling to the bioglass.
  • moieties which provide covalent coupling to the bioglass.
  • bioglass can be functionalised with moieties which enable covalent bonding to the polymer component.
  • One particularly useful method to prepare the hybrid organic-inorganic materials of the invention is via "sol-gel” synthesis. This method is
  • the sol-gel method is efficient in creating a relatively homogeneous distribution between the two phases as a result of dissolving both organic and inorganic compounds in a common solvent.
  • organic molecules are mixed with metal alkoxides and/or organosilanes which undergo sequential hydrolysis and condensation reactions and trap and/or bond to the organic components in the gel structure.
  • TEOS tetraethyl orthosilicate
  • the sol-gel method may be used to prepare polymer-bioglass hybrids.
  • the polymer is PMMA.
  • the sol-gel method may be used to prepare polymer-bioglass hybrids.
  • the polymer is PMMA.
  • the presence of an organosilane coupling agent ⁇ e.g., MPMA copolymerised into the PMMA results in the formation of Si-O-Si covalent bonds to the bioglass silica network, and results in the formation of a relatively homogenous product.
  • organosilane coupling agent e.g., MPMA copolymerised into the PMMA
  • FIG 1 A A conceptual representation of the hybrid material shown in Figure 1 A is shown in Figure IB.
  • the present inventors have utilized analytical methods such as Fourier Transform Infrared (FTIR), 1H, 1 C and 29 Si solid-state NMR spectroscopy to study the microstructure of the hybrids of the invention at the molecular level.
  • FTIR Fourier Transform Infrared
  • 1H 1H
  • 1 C 29 Si solid-state NMR spectroscopy
  • the present invention provides a composite material comprising a covalently coupled polymer-bioglass.
  • the covalently coupled polymer-bioglass is a "hybrid" composite material, meaning that the polymer and bioglass are covalently coupled to each other on the molecular scale.
  • the polymer-bioglass are covalently coupled.
  • other chemical bonds can be utilised to molecularly couple the polymer and bioglass.
  • the constituents of the bioglass are chosen such that the bioglass is adapted to form an apatite layer on the bioglass surface.
  • the hybrid materials of the invention mimic all the properties of bone, are bioactive, osteoconductive and biocompatible.
  • Bioactive glasses of silicate composition which were first developed by Hench and co-workers in 1969 (Hench LL, Splinter RJ, Allen WC, Greenlee TK. Bonding mechanisms at the interface of ceramic prosthetic materials. J. Biomed. Mater. Res. 1971;5(6): 117-141), represent a group of surface reactive materials which are able to bond to bone in physiological environment (Hench LL. Bioceramics. J. Am. Ceram. Soc. 1998;81(7): 1705-1728.).
  • Bioactive glasses suitable for the present invention consist of a silicate network incorporating sodium, calcium and phosphorus in different relative proportions.
  • the classical 45 S5 bioactive glass composition is universally known as
  • Bioglass® (composition in wt%: 45% Si0 2 , 24.5% Na 2 0, 24.5% CaO and 6% P 2 0 5 ) and is a preferred bioglass composition for the present invention.
  • Fabrication techniques for bioactive glasses include both traditional melting methods and sol-gel techniques.
  • the typical feature common to all bioactive glasses, being melt or sol-gel derived, is the ability to interact with living tissue forming strong bonds to bone (and in some cases soft) tissue, a property commonly termed bioreactivity or bioactivity.
  • the bonding to bone is established by the precipitation of a calcium-deficient, carbonated apatite surface layer on the bioactive glass surface when in contact with relevant physiological fluid or during in vivo applications.
  • bioglass is intended to mean those forms of glass which are biologically compatible. Glasses are defined as solid materials with large enormous structural disorder or liquid materials with large viscosity values, whereas bioglasses are considered as a class of bioceramics with extensive applications in biomedical engineering and bone replacement materials. It will be appreciated that there are a number of types of bioglass, and that any bioglass will be suitable for use in the invention. Non-limiting examples of other bioglasses are those which contain other ions such as Zn, P, Mg, Sr, F, Fe, B, K, Na, etc, and combinations thereof.
  • the preferred precursors to the bioglass inorganic component are tetraethyl orthosilicate and tetramethyl orthosilicate, trimethyl orthosilicate, and combinations thereof. However, it will be appreciated by the skilled person that other precursors can be used.
  • a typical composition for a bioglass suitable for the present invention comprises:
  • the polymer is an organic polymer, and is preferably PMMA.
  • PMMA polymethyl methacrylate copolymer
  • other polymers, pre- polymer or oligomers can be utilised in the composite materials of the invention.
  • Suitable polymers have the following polymer compositions:
  • alternating polymerized monomers e.g. monomers A and B polymerized into ABABABAB
  • block copolymers e.g. monomers A and B polymerized into
  • AAAAABBBBBB wherein the block length of each of the blocks is 2, 3, 4, 5 or even more repeat units;
  • graft copolymers consist of a polymeric backbone with side chains attached to the backbone
  • the polymers suitable for the invention may have different polymer architectures including linear, comb/branched, star, dendritic (including dendrimers and hyperbranched polymers).
  • a general review on the architecture of polymers is given by ODIAN, George, Principles of Polymerization, 4th, Wiley-Interscience, 2004. p. 1-18.
  • Comb/branched polymers have side branches of linked monomer molecules protruding from various central branch points along the main polymer chain (at least 3 branch points).
  • Star polymers are branched polymers in which three or more either similar or different linear homopolymers or copolymers are linked together to a single core.
  • Dendritic polymers comprise the classes of dendrimers and hyperbranched polymers. In dendrimers, with well-defined mono-disperse structures, all branch points are used (multi-step synthesis), while hyperbranched polymers have a plurality of branch points and multifunctional branches that lead to further branching with polymer growth (one-step polymerization process).
  • the polymers may be prepared via addition or condensation type polymerizations.
  • Polymerization methods include those described by ODIAN, George, Principles Of Polymerization, 4th edition, Wiley-Interscience, 2004, p. 39-606.
  • Addition polymerization methods include free radical polymerization (FRP) and controlled polymerization techniques.
  • Suitable controlled radical polymerization methods include:
  • Catalytic chain transfer e.g. using cobalt complexes
  • Nitroxide (e.g. TEMPO) mediated polymerizations e.g. TEMPO
  • Suitable examples of monomers for synthesising the polymer component include: acrylic acid, methacrylic acid, maleic acid (or their salts), maleic anhydride, alkyl(meth)acrylates (linear, branched and cycloalkyl) such as methyl(meth)acrylate, n-butyl(meth)acrylate, tert-butyl(meth)acrylate, cyclohexyl(meth)acrylate, and 2-ethylhexyl(meth)acrylate; aryl(meth)acrylates such as benzyl(meth)acrylate, and phenyl(meth)acrylate,
  • hydroxyalkyl(meth)acrylates such as hydroxyethyl(meth)acrylate, and hydroxypropyl(meth)acrylate; (meth)acrylates with other types of
  • oxiranes amino, fluoro, polyethylene oxide, phosphate substituted such as glycidyl (meth) acrylate, dimethylaminoethyl (meth)acrylate, trifluoroethyl acrylate, methoxypolyethyleneglycol (meth)acrylate, and tripropyleneglycol (meth)acrylate phosphate; allyl derivatives such as allyl glycidyl ether; styrenics such as styrene, 4-methylstyrene, 4-hydroxystyrene, 4- acetostyrene, and styrene sulfonic acid: (meth)acrylonitrile; (meth)acrylamides (including N-mono and ⁇ , ⁇ -disubstituted) such as N-benzyl (meth)acrylamide; maleimides such as N-phenyl maleimide; vinyl derivatives such as vinyl alcohol, vinylcaprol
  • Typical condensation type polymers include polyurethanes, polyamides, polycarbonates, polyethers, polyureas, polyimines, polyimides, polyketones, polyester, polysiloxane, phenolformaldehyde, urea-formaldehyde, melamine- formaldehyde, polysulfide, polyacetal or combinations thereof.
  • the initiator can be a thermal initiator or a photo-initiator.
  • Thermal initiator(s) suitable for use in the invention include tertamyl peroxybenzoate, 4,4-azobis(4- cyanovaleric acid), 1, 1' azobis(cyclohexanecarbonitrile), 2,2'- azobisisobutyronitrile (AIBN), benzoyl peroxide, 2.2-bis(tert- butylperoxy)butane, 1, l-bis(tertbutylperoxy)cyclohexane , 1, l-bis(tert- butylperoxy)cyclohexane, 2,5-bis(tert-butylperoxy)-2,5-dimethylhexane, 2,5- bis( tert-butylperoxy)-2,5-dimethyl-3 -hexyne, bis(l-( tert-butylperoxy)
  • Corsslinking monomers can also be utilised.
  • Certain resorbable polymers may be alternatively, or additionally, be used in the invention, such as: polylactides (PLA), poly-L-Iactide (PLLA), poly-DL-Iactide(PDLLA); polyglycolide (PGA); copolymers of glycolide, glycolide/trimethylene carbonate copolymers (PGA/TMC); poly (lactide ethylene oxide fumarate), other copolymers of PLA, such as
  • lactide/tetramethylglycolide copolymers lactide/trimethylene carbonate copolymers, lactide/d-valerolactone copolymers, lactide/s-caprolactone copolymers, L -lactide/D L -lactide copolymers, glycolide/L-Iactide copolymers (PGA/PLLA), polylactide-co-glycolide; terpolymers of PLA, such as lactide/glycolide/trimethylene carbonate terpolymers, lactide/glycolide/ ⁇ - caprolactone terpolymers, PLA/polyethylene oxide copolymers;
  • polydepsipeptides unsymmetrically 3,6-substituted poly-1 ,4-dioxane-2,5- diones; polyhydroxyalkanoates, such as polyhydroxybutyrates (PHB); PHB/b- hydroxyvalerate copolymers (PHB/PHV); poly-b-hydroxypropionate (PHP A); poly-p-dioxanone (PDS); poly-dvalerolactone - poly-8-caprolactone, poly(E- caprolactone-DL-Iactide) copolymers; methylmethacrylate-N-vinyl pyrrolidone copolymers; polyesteramides; polyesters of oxalic acid; polydihydropyrans; polyalkyl-2-cyanoacrylates; polyurethanes (PU); polyvinylalcohol (PVA); polypeptides; poly-b-malic acid (PMLA); poly-b-alkanoic acids;
  • polycarbonates polyorthoesters; polyphosphates; poly(ester anhydrides); and mixtures thereof; and natural polymers, such as sugars, starch, cellulose and cellulose derivatives, polysaccharides, collagen, chitosan, fibrin, hyalyronic acid, polypeptides and proteins.
  • each of the polymers listed above are functionalised with suitable silane functionality adapted to covalently link the bioglass and the polymer components, or include a suitable co- polymerised silane-containing monomer, such as MPMA.
  • non-resorbable polymers may be used in the invention, such as: polymethyl methacrylate, poly methacrylate, poly butyl acrylate, and combinations thereof.
  • polymers listed above are functionalised with suitable silane functionality adapted to covalently link the bioglass and the polymer, or include a suitable co- polymerised silane-containing monomer, such as MPMA.
  • Preferred natural polymers are: collagen, chitosan, fibrin, hyalyronic acid, polypeptides and proteins.
  • Preferred resorbable polymers are: polylactides (PLA), poly-L-Iactide (PLLA), poly-DL-Iactide(PDLLA); polyglycolide (PGA); copolymers of glycolide, glycolide/trimethylene carbonate copolymers (PGA/TMC); poly (lactide ethylene oxide fumarate), copolymers of PL A and polyhydroxyalkanoates.
  • Preferred non-resorbable polymer are: polymethyl methacrylate, poly methacrylate, and poly butyl acrylate.
  • Preferred molecular weights of the polymer are between 2,000 and 1,000,000 g.mol "1 .
  • the polymer component may be functionalised either by introducing functionality to the polymer or cross linking a suitable monomer having the desired functionality.
  • Non-limiting examples of coupling agents are: MPMA, (3-aminopropyl) triethoxysilane, 3- glycidoxypropyldimethoxymethylsilane, acetoxytri-tert- butoxysilane and combinations thereof. Coupling agents with the general formula are also suitable for use with the invention.
  • RO is a hydrolyzable group, such as methoxy, ethoxy,
  • X is an organofunctional group, such as amino
  • the polymer and the bioglass are suitably functionalized such that the functionalisation on each component can be covalently linked to molecularly couple together the polymer and the bioglass.
  • hybrid materials of the invention may be formed into a variety of shapes and used for various medical purposes.
  • a medical device formed from a hybrid of the invention is preferably chosen from the group consisting of: a 3D implantable scaffold, an orthopaedic implant for
  • the hybrid materials of the invention may be implanted into a body in different ways, including, but not limited to subcutaneous implantation, implantation at the surface of the skin, implantation in the oral cavity, use as sutures and other surgical implantation methods. In other embodiments, the hybrid materials of the invention may be injected subcutaneously.
  • the hybrid materials of the invention may be coated with at least one resorbable polymer material, non-limiting examples of which include polyglycolides, polydioxanones,
  • the coating material may comprise healing promoters such as thrombosis inhibitors, fibrinolytic agents, vasodilator substances, anti-inflammatory agents, cell proliferation inhibitors, and inhibitors of matrix elaboration or expression. Examples of such substances are discussed in U.S. Patent No. 6, 162, 537.
  • the present invention also contemplates using a polymer coating, (e.g. a resorbable polymer) in conjunction with a healing promoter to coat the implantable medical device, for example according to the reference [Wu C. Acta
  • the hybrid materials of the invention are a fully synthetic bone graft substitute. Due to its interconnected pores, the material serves as an ideal osteoconductive scaffold and supports the formation of new host bone. As highlighted herein, many of the advantages of the new materials of the invention can be summarised as follows:
  • the uses of the present invention are many fold, including:
  • the hybrid materials of the invention may also be formed into ribbons or fibres.
  • biologically-compatible solvents such as ethanol rather than THF
  • biologically-compatible sodium bicarbonate catalyst rather than hydrochloric acid (HF).
  • hybrids of the invention enables the hybrids of the invention to be utilised as injectable bone cement and interconnected porous scaffolds.
  • the hybrid materials of the invention are mechanically strong and biologically active, which enables them to be utilised as an alternative to currently used bone cements.
  • the hybrid materials of the invention can be used as a porous tissue engineering scaffold.
  • the scaffold has a porosity of at least 60 %, more preferably at least 80 %, and most preferably at least 90 %.
  • the hybrid of the invention may be formulated and prepared to have an interconnected porosity of 20, 25, 30, 35, 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, 90 or 95%.
  • the pore sizes are between about 75 to about 300 ⁇ .
  • the hybrid materials of the invention could be configured to have lower or greater pore size according to the intended or desired use, and any pore size would be within the purview of the present invention.
  • the compressive strength of the hybrid materials of the invention are preferably between about 2 to 20 MPa, but can be greater than this depending on the choice of polymer and bioglass, and their relative concentrations.
  • the hybrid materials of the invention have a biocompatibility when placed in physiological fluid.
  • the hybrid materials of the invention form a hydroxyapatite layer upon exposure to bodily fluids.
  • the formation of hydroxyapatite is widely recognised as strong evidence that the body accepts the material as sui generis and is a requirement for the implant to chemically bond with living bone and tissue.
  • the material of the invention has favourable cell-interaction properties. For example, two types of osteoblast cells (primary human osteoblast cells and mouse clonal osteoblast cells) adhere to and proliferate on the hybrid material.
  • the material of the invention induces mineralization for cultured cells.
  • the hybrid material of the invention has homogenous morphological features.
  • the hybrid material of the invention has hardness similar to the hardness of natural bone.
  • the hybrid material of the invention shows no cytotoxicity in in vivo tests.
  • the present invention provides a method of production of a composite material comprising the steps of:
  • the inorganic sol of bioglass precursors is selected from tetraethyl orthosilicate ("TEOS”), trimethyl orthosilicate, and combinations thereof.
  • TEOS tetraethyl orthosilicate
  • trimethyl orthosilicate and combinations thereof.
  • the polymer is a sol of poly(methyl methacrylate)-co-3-
  • PMMA-co-MPMA (trimethoxysilyl)propyl methacrylate
  • the polymer may by resorbable, or non-resorbable, and can be chosen from the polymers listed above.
  • bioglass precursors and polymer(s) are dissolved or suspended in the same solvent.
  • Preferred solvents are selected from:
  • TFIF tetrahydrofuran
  • MP N-methyl-2-pyrrolidone
  • acetic acid tetrahydrofuran
  • the inorganic sol is prepared by hydrolysing TEOS in low pH aqueous solution (pH between 1 to 6) and with the addition of calcium chloride or any other suitable calcium salt at room temperature.
  • the aqueous solution can be acidified with hydrochloric acid.
  • the TEOS:water molar ratio is between 1 :4 to 1 :8, for example 1 :5, 1 :6, or 1 :7.
  • the polymer is not greater than 80 vol.%.
  • the ratio of polymer to bioglass precursors is between 1 : 1 to 1 :6.
  • the ratio is 1 : 1, or 1 :2, or 1 :3, or 1 :4, or 1 :5 or 1 :6.
  • the reaction can proceed between couple of seconds up to several minutes. However, the gelation and drying procedure may take from hours to days.
  • the PMMA-co-MPMA is prepared from free radical polymerisation of PMMA and MPMA in a predetermined molar ratio, using ⁇ , ⁇ '-azoisobutyronitrile ("AIBN") as an initiator.
  • the predetermined molar ratio of MPMA to PMMA is most preferably between about 0.1 to 0.5 molar ratio.
  • the molar ratio of MPMA:MMA is 0.1 to avoid undesired crosslinking between the copolymer chains.
  • a ratio lower ratio than 0.1 results in reduced, poor or a lack of formation of hybrid Si-C bonds in the resulting composite material.
  • the polymerization reaction can take anywhere from 1 to 12 hours and be conducted at temperatures between 60 to 100°C.
  • the sols are at least partially gelled. In other embodiments the hybrid material is dried and sintered.
  • an organic-inorganic hybrid composite when synthesised by a method according to the second aspect.
  • the present invention provides a formulation for preparing a composite material, the formulation comprising a bioglass sol, a polymer sol having functionality adapted to covalently couple the bioglass and polymer.
  • the polymer is a copolymer of MPMA:MMA at a 0.1 molar ratio, and wherein the polymer composition is above 60 vol. %.
  • the present invention provides a kit comprising in separate containers: a bioglass sol, a copolymer sol and a catalyst.
  • the covalent coupling between the bioglass and copolymer sols is catalysable by the addition of sodium bicarbonate to produce a solidified hybrid composite material.
  • Other catalysts will be known to the skilled person.
  • the bioglass sol and copolymer sol are in a separate containers, and in other embodiments are in the same container.
  • the invention provides a method of forming a composite material in situ or in vivo, the method comprising the steps of delivering to a subject in need thereof a polymer sol and a bioglass sol to a predetermined location in the body of the subject and in situ catalysing the formation of a hybrid bioglass-polymer composite material.
  • the predetermined location can be subcutaneous, for example, at a site in the body in need of a scaffold for bone growth, such as a region of bone damage or weakness.
  • the polymer sol and bioglass sol can be administered sequentially or preferably simultaneously.
  • the method preferably comprises mixing the bioglass sol and copolymer sol at elevated temperature.
  • the elevated temperature is between 70 and 95°C, and is preferably 70°C.
  • the mixing time can be seconds to hours, and is preferably 30 seconds.
  • the heated solution is then cooled to room temperature and the catalyst solution added.
  • the resulting catalysed liquid is then drawn into a syringe and injected subcutaneously into the defected site.
  • the method preferably comprises mixing bioglass and polymer sols at elevated temperature obtain a homogenous and well-dispersed solution.
  • the elevated temperature is between 50 and 95°C, and is preferably 70°C.
  • the mixing time can be seconds to hours, and is preferably 30 seconds.
  • the hybrid is then cooled to room temperature and sodium bicarbonate is added a foaming agent to create porosity within the scaffold. Other foaming agents are well known to the skilled person.
  • the temperature of the solution is then elevated to form bubbles from degradation of sodium bicarbonate and to stabilize the porous structure during gelation of hybrid composite.
  • the elevated temperature is between 50 and 95°C, and is preferably 70°C.
  • the resulting porous scaffold (hybrid bioglass-polymer composite) is then dried, preferably at room temperature.
  • the present invention provides the use of a hybrid composite of the invention as a bone filler, as an injectable formulation, or as a porous scaffold for biomedical purposes.
  • the hybrid of the invention is shaped into the desired application by controlling the gelation time using process parameters such as temperature and solvent and also, the catalyst (sodium bicarbonate) for the condensation reaction of compounds.
  • the composite material forms a hydroxyapatite layer upon exposure to bodily fluids.
  • An implantable medical device comprising the composite material of the invention.
  • the medical device is formed into a device chosen from: a 3D implantable scaffold, an orthopaedic implant for reconstructive surgery, a dental implant/prostheses, a spine implant, implants for craniofacial
  • the medical device may be permanently implanted or temporarily implanted.
  • the medical device is substantially biodegradable.
  • the medical device has a porosity of between about 10 to about 80%, a pore size is between about 20 to about 500 micron.
  • the medical device is coated with at least one resorbable polymer material selected from polyglycolides, polydioxanones, polyhydroxyalkanoates, polylactides, alginates, collagens, chitosans, polyalkylene oxalate,
  • the medical device is coated with at least one healing promoter selected from thrombosis inhibitors, fibrinolytic agents, vasodilator substances, anti-inflammatory agents, cell proliferation inhibitors, and inhibitors of matrix elaboration or expression.
  • Bone implant or biocement comprising the hybrid composite of the invention.
  • a method for producing an implantable medical device comprising: transferring the hybrid composite of the invention onto a substrate thereby forming said implantable medical device.
  • An implantable drug delivery device comprising the hybrid composite of the invention.
  • a method for improving the long term stability of an implantable medical device comprising the step of: coating said device with the hybrid composite of the invention.
  • hybrid composite of the invention in the regeneration or resurfacing of tissue, comprising contacting the tissue with a quantity of the hybrid composite of the invention for a sufficient period to at least partially effect said regeneration or resurfacing.
  • a method for regenerating or resurfacing tissue comprising the step of: contacting said tissue with the hybrid composite of the invention.
  • a method for forming osseous tissue on an orthopaedic defect comprising the step of: contacting said defect with the hybrid composite of the invention.
  • Figure 1 A is a schematic molecular structure of a copolymer of PMMA- co-MPMA-bioglass class II hybrid material
  • Figure IB is a schematic representation of Fig. 1 A
  • Figure 2 show the 400 MHz 1H NMR spectra of a) P Co nt, b) PLOW, c) P Co nt expansion in the region of 3.5-4.1 ppm, and d) P LOW expansion in the region of 3.5-4.1 ppm in deuteriochloroform (CDC1 3 ).
  • Figure 3 is the 1 C NMR spectrum of the P Me d PMMA-co-MPMA copolymer of the present invention.
  • the methyl and carbonyl carbons have been assigned as indicated.
  • the spectrum included the characteristic peaks of carboxyl group (178 ppm), methylene (56 ppm), methoxy (52 ppm), quaternary (45 ppm) and methyl carbons (17 ppm).
  • Figure 5 depicts optical images of hybrid samples a) Hc on t6o, b) H H i g h6o and c) Hcont4o- It was observed that by increasing the molar ratio of MPMA in the copolymers structure, firmer and more transparent monoliths were formed;
  • Figure 6 is the 1 C CPMAS NMR spectra of a) neat PMMA (P Co nt), b)
  • PMMA with high coupling agent (PHigh) and c) H H i g h6o hybrid These spectra were acquired with high power 1H decoupling.
  • the -0-CH 2 - peaks of the coupling agent MPMA chains were also visible at 67 ppm;
  • FIG. 7(a) is the simplified schematic structure of hybrid in the absence of Ca 2+ .
  • Q and T are parameters that describe the hybrid network connectivity.
  • Q n shows a silicon bonds to n other silicons via "bridging" oxygen atoms, and the T sites indicate a silicon atom bonds to one carbon atom (Si-C) and to other silicons;
  • Figure 7(b) shows the 29 Si CPMAS NMR spectra of hybrid samples set against bioglass.
  • the Q- and T-site silicons are assigned as indicated. Three main resonances are observed at -110, -100 and -90 ppm that correspond to the Q 4 , Q 3 and Q 2 species, respectively.
  • Figure 8(a) is a Directly-Polarised 29 Si HPDMAS NMR of PMMA- bioglass hybrids. The acquired spectra are shown in bold while deconvoluted peaks fits are shown as thin lines. The Si peaks were deconvoluted to quantify the bioglass structure in the composites;
  • Figure 8(b) shows the relative populations for the different Q sites in the hybrid materials.
  • the relative amount of Q 4 species was c. 15-22 % higher as compared to hybrids with high coupling agent (H H i g h6o and H H i g h8o), while hybrids with low and medium coupling agents (H LoW 6o and H Me d6o) possessed intermediate concentrations of Q 4 species.
  • Figure 9 shows the 2D 1H- 1 C and 1H- 29 Si HetCor for H H i gh 6o.
  • the ID 1H projections are plotted to the left of the 2D spectra while the 1 C and 29 Si projections are plotted at the top.
  • Figure 10 shows the FTIR spectra for a) Hc on t6o, b) H H i g h6o, c) Pcont, and d) bioglass.
  • the peaks at 1030-1040, 930 and 790 cm “1 corresponded to Si-O-Si network structure and were detected only for bioglass and H H i g h6o-
  • the peaks at 1625 cm “1 and 3370 cm “1 correspond to the presence of C0 3 2" and OH groups, respectively, of the bioglass composition.
  • FIG 11 shows SEM images for a) H Co nt6o and b) H ffig h6o.
  • the H Co nt6o morphology was non-homogeneous and the phase separation between the bioglass and polymer was visible on the micron scale.
  • the chemically-coupled hybrid H H i g h6o
  • Figure 12 shows the TGA (a, top) and DTG (b, bottom) curves of a) PMed, b) bioglass, c) H Co nt6o, d) H Me d6o, e) H H i g h6o.
  • TGA profile of neat PMMA-co-MPMA copolymer one observes three distinct peaks at 165 °C, 270 °C and 360 °C. These peaks are attributed to the presence of head-to-head linkages, end-chain unsaturation and random scission within the polymer chain, respectively.
  • DTG profile of bioglass a single peak at 100 °C is due to the loss of water;
  • Figure 13 shows the DSC curves of a) Pcont, b) Pffigh, c) pure bioglass, d) Hcont6o, e) H H i g h6o-
  • the peak at 120 °C which was observed in bioglass DSC profile was attributed to the condensation of bioglass silica network.
  • the same trend was also observed in Hc on t6o and H H i g h6o curves.
  • this peak was shifted toward higher temperatures (i.e., to 150 °C) in H H i g h6o due to the covalent bonding of polymer chains to silica network.
  • inorganic moieties reduced the polymer chains mobility and the covalent bonds between the phases reduced the free movement of both organic and inorganic chains.
  • more energy (heat) was required in hybrid samples to approach the same degree of freedom. It could then be concluded that covalent bonding of silica to PMMA chains may have enhanced the thermal stability of the final product.
  • Figure 14 is a table showing optical transparency and gelation time of various samples.
  • Figure 15 shows SEM images of a) bioglass, b) HO, c) HI . Phase separation between ceramic and polymer was evident for the physical mixture of PMMA and bioglass (HO). However, the absence of phase separation and homogenous surface structure for hybrid HI underlined the chemical conjugation and structural integrity of these samples;
  • Figure 16 shows the EDS results of a) HO and b) HI including the EDS mapping results for carbon, silicon and calcium elements and the statistical results of the presence of elements. Three random points were selected in each sample and the average amount of calcium was quantified. The results demonstrate that while the distribution of calcium in HO was dramatically changed, it possessed homogeneous distribution with narrow range variation for HI;
  • Figure 17 shows STEM images of a) HO and b) HI, with a 1 ⁇ scale ruler shown at the bottom left hand corner of each image.
  • HO physical mixture
  • HI acquired from covalent bonding was homogenous even in the nanoscale range.
  • FIG 18 shows the Atomic Force Microscopy (AFM) results of hybrids including topographic images, phase imaging and roughness (Ra) of the surface of the HO and HI for 5 x 5 ⁇ area.
  • AFM Atomic Force Microscopy
  • Figure 19 relates to the mechanical properties of bioglass and HI samples, a) A stress-strain diagram, b) a microhardness analysis (p ⁇ 0.001).
  • the Young's (elastic) modulus of the HI hybrid obtained from the stress-strain diagram is 40-times higher compared to the bioglass samples;
  • Figure 20 shows images of a) bioglass compared to b) the HI hybrid after free-fall from 45 cm height on a wooden surface. The formation of cracks in the structure of the bioglass was clearly visible, but the integrity of hybrid HI was maintained;
  • Figure 21 depicts the bioactivity of the a) bioglass, b) HI samples after 7 days incubation in SBF at 37°C. A 10 ⁇ scale ruler is shown to the bottom left hand corner of each image. The results show that the apatite layer was formed on the surface of both bioglass and hybrid.
  • Figure 22 shows SEM images of (a, b) pure PMMA (P 0 ); (c, d) HO; (e, f) HI; and (g, h) bioglass.
  • Both the physical mixture (HO) and chemically conjugated samples (HI) supported cell attachment. However, this attachment was enhanced significantly in HI .
  • a homogenous filopodia cell migration was observed on the surface of HI, which promotes cells anchoring to the biomaterial surface and enhances cell spreading.
  • cells were spread well on the surface and base lamellipodia were well- developed;
  • Figure 23 is an MTS assay of the HI (H H i g h6o) hybrid sample compared with the PMMA and bioglass. As shown, a significant difference ( ⁇ 0.001) was observed in the absorbance of hybrid samples (HI) compared to the pure PMMA and bioglass.
  • PMMA and bioglass had a positive impact on cell interaction and provide a superior environment for cell adhesion and proliferation
  • Figure 24 is a degradation profile of the a) bioglass, b) HI sample.
  • Dashed lines indicate the linear model fitted to the curve. It was observed that the presence of PMMA covalently bonded to the bioglass decreased the rate of degradation of the bioglass. As shown, the degradation rate of the hybrid sample was 1.5 times less than the bioglass.
  • Figure 25 are AFM results including topographic images, phase imaging and roughness of the surface of the HO (physical mixture) and HI (hybrid).
  • Figure 26 are the mechanical properties of bioglass and HI samples, a)
  • Figure 27 shows degradation profiles of hybrid compared with pure PMMA and bioglass (The data was extrapolated for BG until day 83).
  • Figure 28 shows FTIR spectra of a) HI, b) residues of degraded HI.
  • Figure 29 are bioactivity test shows the formation of HA on the surface of a) bioglass, b) HI, c) HO and d) PMMA samples after 7 days incubation in SBF.
  • Figure 30 are SEM images of a) Pure PMMA, b) HO, c) HI, d) bioglass. White arrows show the individual narrowed cells on the surface of samples.
  • Figure 31 is MTS assay of HI hybrid sample compared to the PMMA and bioglass.
  • Figure 32 shows ALP staining on materials showing ALP+ cells (blue) after 4 days of osteogenic differentiation. Positive staining was seen on HI hybrid and tissue culture plastic, but not on PMMA, bioglass, or HO material.
  • Figure 33 shows tissue reaction to the bioglass implant at day 10 after implantation, a) Azan- staining, the vessel-rich (arrows) multi-layered tissue wall (double arrow) covering the bulk-like structural segments of the implanted material (BG). A layer of mononuclear cells (squares) was adherent to the material surface, b) Macrophage-specific F4/80-immunostaining, macrophages (pentagons) were located in all regions of the peri-implant tissue (double arrow), c) TRAP-staining, only a low number of mononuclear cells within the peri-implant tissue expressed TRAP (diamonds), while the majority of these cells were TRAP-negative (black arrow heads).
  • Figure 34 shows tissue reaction to bioglass fragments at day 10 after implantation, a) H&E-staining, an overview of the implantation bed of fragments (BG), which were embedded within a vessel-rich (arrows) granulation tissue. At their surfaces mononuclear (pentagons) and
  • Figure 36 shows the optical transparency of PMMA-co-MPMA
  • Figure 37 shows the effect of temperature and solvent on the gelation time of PMMA-bioglass hybrids.
  • Figure 38 shows the effect of concentration of SB on the gelation time of HlEtOH at room temperature.
  • Figure 39 shows the effect of gelation time and network structure on the mechanical properties of hybrids, a) stress-strain curves, b) toughness value, c) ultimate stress, and d) ultimate strain.
  • Figure 40 are SEM images of hybrids after 7 days soaking in SBF, a) HI, b) HlEtOH, c) EDS analysis of crystals formed on the surface of hybrids.
  • Figure 41 show injectability of the hybrid solution (HlEtOH) by addition of SB.
  • Figure 42 provides the degradation profiles of HI and HlEtOH for the period of 100 days.
  • Figure 43 shows images of porous PMMA-bioglass scaffold, a) SEM micrograph, b) ⁇ -CT scan of PMMA-bioglass porous scaffold in different cross- sections and the ratio of closed and open pores.
  • Figure 44 shows the mechanical properties of HlEtOH monoliths and porous structures - a) stress-strain curves, and d) ultimate strain.
  • Figure 45 provides the viability assay of HlEtOH samples in the form of monoliths and porous scaffolds (PS).
  • Figure 46 provides the alizarin red S staining images of monoliths and porous scaffolds.
  • Figure 47 provides the quantified analysis of mineralization study of HIEtOH samples in the form of monoliths and porous scaffolds (PS).
  • the terms 'predominantly' and 'substantially' as used herein shall mean comprising more than 50% by weight, unless otherwise indicated.
  • the recitation of a numerical range using endpoints includes all numbers subsumed within that range (e.g., 1 to 5 includes 1, 1.5, 2, 2.75, 3, 3.80, 4, 5, etc.).
  • the terms 'preferred' and 'preferably' refer to embodiments of the invention that may afford certain benefits, under certain circumstances.
  • bioactive material a material that has been designed to elicit or modulate biological activity.
  • Bioactive material is often surface-active material that is able to chemically bond with the mammalian tissues.
  • a biodegradable material is a material that breaks down in vivo, but with no proof of its elimination from body.
  • bioresorbable in this context means that the material is disintegrated, i.e. decomposed, upon prolonged implantation when inserted into mammalian body and when it comes into contact with a physiological environment.
  • the by-products of a bioresorbable material are eliminated through natural pathways either because of simple filtration or after their metabolisation.
  • bioresorbable and resorbable can be used
  • resorbable glass means silica-rich glass that does not form a hydroxyl-carbonate apatite layer on its surface when in contact with a physiological environment. Resorbable glass disappears from the body through resorption and does not significantly activate cells or cell growth during its decomposition process.
  • bioabsorbable it is meant a material that can dissolve in body fluids without any molecular degradation, and then excreted from the body.
  • biomaterial a material intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ or function of the body.
  • biocompatibility is meant the ability of a material used in a medical device to perform safely and adequately by causing an appropriate host response in a specific location, causing no foreign-body reactions and being non-toxic.
  • resorption is meant decomposition of biomaterial because of simple dissolution.
  • composite is meant a material comprising at least two different constituents, for example a polymer and a ceramic material, such as glass.
  • an implant according to the present context comprises any kind of implant used for surgical musculoskeletal applications such as screws, plates, pins, tacks or nails for the fixation of bone fractures and/or osteotomies to immobilize the bone fragments for healing; suture anchors, tacks, screws, bolts, nails, clamps, stents and other devices for soft tissue-to-bone, soft tissue-into-bone and soft tissue-to-soft tissue fixation; as well as devices used for supporting tissue or bone healing or regeneration; or cervical wedges and lumbar cages and plates and screws for vertebral fusion and other operations in spinal surgery.
  • Precursors required for the synthesis of PMMA-co-MPMA copolymer including MPMA, ⁇ , ⁇ '-Azoisobutyronitrile (“AIBN”) and ⁇ , ⁇ '- dimethylformamide (“DMF”) were purchased from Sigma- Aldrich and used as received.
  • TEOS was mixed with deionised water and HC1 and stirred for 30 minutes followed by addition of calcium chloride dihydrate and stirred for another 30 minutes to yield a homogenously-mixed solution.
  • a common calcium source for the preparation of sol-gel derived bioglasses is calcium nitrate tetrahydrate.
  • calcium chloride was used to minimise the risk of toxicity resulted from the nitrate byproduct.
  • the solutions were then immediately cooled down to room temperature by water bath.
  • Sodium bicarbonate (SB) was added as a solution in miliQ water with a specific concentration to the hybrid solution. Gelation occurred at room temperature and dried for 24 hours at ambient temperature unsealed. It should be noted that fast gelation did not damage the monolithic structure of these samples.
  • the prepared gel was characterised by solid state NMR analyses to determine the structure of solid network. All the NMR experiments were carried out on a Bruker Biospin Avance III solids-300 MHz and Avance II standard bore, 700 MHz spectrometers (Bruker-Biospin, Rheinstetten, Germany). A Bruker 4-mm double resonance magic-angle spinning ("MAS") probehead was used with MAS frequencies of 5 kHz for 1 H- 29 Si HetCor and 12 kHz for other experiments. The 90 degree pulse length was 4 for 1H and 4.5 for 29 Si and 1 C nuclei.
  • MAS double resonance magic-angle spinning
  • a 300 second recycle delay was used in direct-polarisation 29 Si MR experiments.
  • the SPINAL-64 1 H- 29 Si and 1H- 1 C heteronuclear decoupling strength was 80 kHz during 29 Si and 1 C detection.
  • Hartman-Hahn cross- polarisation (HHCP) was employed for experiments requiring polarisation transfer from 1H to 1 C or 29 Si with a contact times of 1 ms and 4 ms, respectively.
  • 1 H- 29 Si and 1H- 1 C HetCor experiments were carried out with frequency- switched Lee-Goldburg (FSLG) for 1H-1H homonuclear decoupling and 1H chemical shifts were scaled by 0.571 accordingly during data processing. Also attempted were 160 ti increments of 51 ⁇ .
  • FSLG frequency- switched Lee-Goldburg
  • the synthesised copolymers were characterised for the weight average molecular weight (Mw), number average molecular weight (Mn) and
  • polydispersity index (Mw/Mn) by GPC with a Shimadzu Prominence HPLC (Kyoto, Japan) equipped with a GPC column (Jordi Gel DVB mixed bed, Alltech) diode array detector (SPD-M20A, Shimadzu) and refractive index detector (RID- 1 OA, Shimadzu).
  • PDI polydispersity index
  • a TR-FTIR Attenuated total reflection Fourier transform infrared
  • the surface microstructures and phases formed on the specimens were scanned by field emission scanning electron microscopy (FE-SEM; Zeiss ULTRA plus). Samples were mounted on aluminium stubs using conductive carbon paint, then gold coated by using Emitech K7550X instrument prior to SEM analysis.
  • FE-SEM field emission scanning electron microscopy
  • FE- SEM field emission scanning electron microscopy
  • Zeiss ULTRA plus This instrument was equipped with Bruker XFlash 4010 EDS detector with high speed acquisition and hypermapping capability. Samples were mounted on aluminium stubs using conductive carbon paint, then gold coated by using Emitech K7550X instrument prior to SEM analysis.
  • STEM analyses were conducted to investigate the interaction between the phases in nanoscale. Powders were embedded in epoxy resins and microtomed with Leica Ultracut ultramicrotomes (UC7) for 100 nm layers. The layers were harvested and seated on carbon grids prior to STEM analysis (Zeiss ULTRA plus).
  • UC7 Leica Ultracut ultramicrotomes
  • Thermogravimetric analysis was carried out to study the thermal decomposition of the various samples. Analyses were conducted over the temperature range of 30 °C to 600 °C at heating rate of 20 0 C.min _1 under a nitrogen atmosphere. 2.8 Differential scanning calorimetr (DSC)
  • the thermal properties of the polymers were analysed by DSC (TA Q- 1000, USA). The samples had an average weight of 4 mg and were heated from room temperature to 400 °C under constant nitrogen flow (50 mL.min "1 ).
  • the surface characteristics of the prepared hybrid samples were investigated by atomic force microscope (AFM; Asylum Research, MFP-3D- BIO) in AC mode. Phase imaging was used to identify different phases. A thin film was required to examine the surface properties.
  • PMMA-co-MPMA copolymers were synthesised with MPMA:PMMA molar ratios of 0, 0.004, 0.02 and 0.1 which were coded as control polymer (Pcont), polymers with low (PLOW), medium (PMed) and high (Pffigh) ratio of functional groups, respectively (see, Table 2).
  • the suffix defines the ratio of functionalisation of a copolymer that corresponds to the ability for covalent bond formation with bioglass.
  • Fig.2(a,b) shows the 1H NMR results for Pc ont and P LOW in deuteriochloroform (CDC1 3 ). The presence of peaks at 3.59 ppm (a, - OCH3), 1.014 ppm (b, -CH 2 ) and 1.621 (c, -CC3 ⁇ 4) in 1H NMR spectrum confirmed the complete synthesis of Pc 0 nt PMMA).
  • the 1 C NMR spectrum of PMMA-co-MPMA copolymers synthesised in this study is provided in Fig.3.
  • the 1 C NMR spectrum of PMed included the characteristic resonances of carboxyl group (178 ppm), methylene (56 ppm), methoxy (52 ppm), quaternary (45 ppm) and methyl carbons (17 ppm). These data confirm the successful copolymerisation of PMMA and MPMA. A similar spectrum was obtained for P Low and Pmgh samples.
  • Bioglass (sol (A)) having a composition of TEOS : water : HC1 :
  • Hc on t4o is the hybrid sample prepared from Pc on t polymer and the H 40 mixture of sol (B) : sol (A) (i.e., 40 : 60). These hybrids were examined visually, to determine the effect of composition of polymer solution to bioglass solution and the fraction of MPMA on the transparency, phase separation and gelation behaviour of samples.
  • Transparency is an indicative parameter for the absence of phase separation below the scale of 400 nm (see, e.g., Wei, et al., J. Appl. Polym. Sci., 1998, 70, 1689-1699).
  • the samples that had their transparency and gelation time examined are listed in Table 3. It was observed that by increasing the molar ratio of MPMA in the copolymers, firmer and more transparent monoliths were formed. However, at low MPMA : PMMA molar ratio (e.g., ⁇ 0.1) samples were less transparent and gelation time was greater than 20 days at ambient temperature (see, Fig.2(a,b)).
  • the composition of the bioglass (sol (A)) was another factor that elicited a significant impact on the phase separation, transparency and gelation time of the hybrids.
  • a one-phase solution was acquired for Hc on t6o and H H i g h6o-
  • high proportion of bioglass e.g., Hc on t4o >50 vol.%
  • two separate phases were formed due to the agglomeration of the polymer chains (see, Fig.5(c)).
  • NT not transparent
  • NG no gel formation before 20 days
  • the gelation time of the H H i g h samples was tuned by varying the composition of the bioglass.
  • the gelation time of the hybrid material was decreased from 20 days to about 2 hours at room temperature when the composition of polymer solution was increased from 40 vol.% to 80 vol.%.
  • microhardness identification number for a hybrid sample containing the highest MPMA molar ratio was more than 15 times higher than the pure bioglass. This result shows that addition of covalent bonding between organic-inorganic compounds not only accelerates the gel formation, but also dramatically enhances the mechanical properties of the resultant hybrids.
  • Fig.7(a) The schematic illustration of the molecular structure of the proposed hybrid is shown in Fig.7(a).
  • Q and T are parameters that describe the hybrid network connectivity.
  • Q n shows silicon bonds to n other silicons via "bridging" oxygen atoms, and the T sites indicate a silicon atom bonding to one carbon atom (Si-C) and to other silicons ⁇ see, Mahony, et al., Adv. Funct.
  • T 3 shows that a silicon atom bonds to three silicon and one carbon atom.
  • Three main peaks are observed at -110, -100 and -90 ppm in the qualitative 29 Si CPMAS NMR in Fig.7(b), that correspond to the Q 4 , Q 3 and Q 2 species, respectively.
  • T n sites are only visible in the H H i g h6o and H H i g h8o hybrids at -50, -60 and -65 ppm for T 1 , T 2 and T 3 , respectively.
  • the presence of these peaks confirms the formation of Si-C bonds - and hence, demonstrates the contribution of the polymer into the bioglass network structure.
  • hybrids with high coupling agent i.e., H H i g h6o and H H H i g h8o
  • hybrids with low and medium coupling agents i.e., H LoW 6o and H Me d6o
  • the 2D 1H- 1 C and 1H- 29 Si HetCor NMR spectra of the H H i gh 6o material conclusively identified the nanometer-scale interaction between the polymer and bioglass phases.
  • the 1H chemical shifts correlating to 1 C- 29 Si chemical shifts were spread in a 2D map, with the intensities plotted as contour levels as shown in Fig.9.
  • the 2D 1 H- 1 C HetCor was acquired with a short contact time of 100 of cross-polarisation transfer from the 1H to the 1 C nuclei.
  • this spectrum primarily shows the localised correlations of 1 C species to their directly bonded 1H species.
  • FTIR was also used to investigate the molecular structure of the prepared hybrids.
  • the FTIR spectra of neat bioglass, PMMA ⁇ i.e., Pcont), hybrid of bioglass and highly functionalised copolymer ⁇ i.e., H H i g h6o), and the physical mixture of bioglass and polymer ⁇ i.e., Hc on t6o) are shown in Fig.10 ⁇ i.e., (a) Hconteo, (b) H H igh6o, (c) Pcont, and (d) bioglass).
  • the peaks at 1030-1040 cm “1 , 930 cm “1 and 790 cm “1 corresponded to Si-O-Si network structure and were detected only for bioglass and H H i g h6o- These peaks were not observed in Hc on t6o due to the lack of gel formation in this sample.
  • the peaks at 1625 cm “1 and 3370 cm “1 correspond to the presence of C0 3 2" and OH groups, respectively, of the bioglass composition.
  • the FTIR results show that the amount of MPMA coupling agent played a critical role on the molecular interactions of compounds from weak to strong bonding that may result in the production of materials with unique physical properties such as morphological features and thermal stability.
  • Hc on t6o and H H i g h6o hybrids were investigated by scanning electron microscopy (SEM). As shown in Fig. 11(a), the Hc on t6o morphology was non-homogeneous and the phase separation between the bioglass and polymer was visible on the micron scale. However, in Fig. 11(b), for the chemically-coupled hybrid (H H i g h6o) a relatively homogenous and smooth surface was detected without any indication of large-scale phase separation. These images confirm that phase separation may be avoided by the formation of covalent bonding between PMMA and bioglass phases. 3.5 Thermal behaviour
  • the thermal behaviour of the synthesised material is an important characteristic from the standpoint of its potential applications - especially when a high temperature process is required, for example, in a sterilisation processes for bioimplants. It was anticipated that the nanoscale interaction between bioglass and polymer would impact significantly upon the characteristics of a hybrid material. As shown in Fig. 12(a), the TGA profile of neat PMMA-co- MPMA copolymer showed three distinct peaks at 165 °C, 270 °C and 360 °C. These peaks are attributed to the presence of head-to-head linkages, end-chain unsaturation and random scission within the polymer chain, respectively. The polymer was degraded completely by 400 °C due to the breakdown of its backbone structure.
  • the TGA profile of bioglass depicted in Fig. 12(b) shows a single peak at 100 °C, which is due to the loss of water.
  • a steady state profile was observed for bioglass up to 600 °C because of thermal stability of the silica structure. .
  • the onset thermal decomposition temperature ("OTDT") of Hc on t6o was increased compared with that of the pure polymer due to the addition of silica, as listed in Table 4.
  • the formation of covalent bonding between organic and inorganic components had the same impact upon the thermal characteristics of the H Me d6o and H H i g h6o hybrids.
  • the degradation peaks were at 360 °C, 381 °C, 390 °C and 400 °C for P Me d, H Co nt6o, H Me d6o, H H i g h6o, respectively.
  • the OTDT of these samples was tuned to higher temperatures by increasing the MPMA composition which greatly impacted the molecular structure and level of chemical interaction.
  • Hc on t6o and H Me d6o possessed three major distinct weight losses in TGA and DTG curves, which were in agreement with the degradation profile of the synthesised polymer.
  • weight losses were detected only at two distinct temperatures (249 °C and 400 °C). This is thought to have occurred due to the presence of strong covalent bonds between PMMA chains and silica structure that inhibit thermal degradation within the range of 200 °C to 300 °C.
  • the OTDT of H H i g h6o was also higher than other hybrid products having lower levels of coupling, such as H Me d6o and H Low 6o.
  • the degradation profile of the hybrid was also influenced by the presence of covalent bonding. A similar trend was observed for the degradation profile of the hybrid H con t6o ⁇ see, Fig.12A(c) and Fig.12B(c)) and the polymer (P Med ) ⁇ see, Fig.12A(a) and Fig.12B(a)) at temperatures below OTDT of the polymer ⁇ i.e., 360 °C).
  • Fig.13(a,b) show that the glass-transition temperature (T g ) of pure PMMA was 105 °C, whereas no thermal effect was observed in DSC profile of highly functionalised PMMA (P H i gh ) below 250 °C. This effect might be due to the covalent bonding of MPMA-PMMA that limited the mobility of PMMA polymer chains. This effect decreases the degree of freedom of the PMMA chains and results in the absence of a T g peak for the P ff ig h copolymer.
  • Table 4 Peaks of the thermal decomposition of samples and their S1O 2 content
  • Selected samples were characterised for their bioactivity by incubation in simulated body fluid ("SBF") at 37 °C ⁇ i.e., body temperature) for the periods of 1, 3, 7, and 14 days. After such time, the samples were washed twice with deionised water to remove any residue of minerals absorbed on the surface and dried in the oven at 37 °C for 24 hours. The dried monoliths were examined to determine their morphological characteristics and investigate the formation of any apatite layer on the surface.
  • SBF simulated body fluid
  • HOB Primary human osteoblast
  • the monolith samples (each, 12 mm diameter x 1 mm height) were placed into a well plate and kept in 70% ethanol for one hour to sterilise followed by rinsing with phosphate buffered saline ("PBS"), triple. Samples were then exposed to UV light for 30 minutes and were washed with fresh medium at 37 °C overnight. The substrates were placed into well tissue- culturing polystyrene plates and cells with the density of 2 ⁇ 10 5 cells/mL were seeded onto the monoliths and kept in a C0 2 incubator at 37 °C for four days.
  • PBS phosphate buffered saline
  • Osteoblast cell proliferation was determined by the 3-(4,5- dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H- tetrazolium (MTS) assay.
  • MTS tetrazolium
  • hybrid samples were prepared by the sol gel method in which an aqueous solution of bioglass (sol (A)) was mixed with a solution of PMMA (sol (B)) at 70 °C.
  • samples containing MPMA are coded as "HI” ⁇ i.e., “hybrids" and those without coupling agents represent the physical mixture and are coded as "HO”. Pure bioglass was also considered as the control sample.
  • Fig.15 The surface morphology of neat bioglass, HO and HI hybrids were compared in Fig.15. Phase separation between the ceramic and polymer was evident in Fig.15(b) for the physical mixture of PMMA and bioglass (HO). However, the absence of phase separation and the presence of a relatively homogenous surface structure in Fig.15(c) for the hybrid (HI) underlined the chemical conjugation and structural integrity of these samples. This result was further confirmed by mapping the SEM image by EDS detector for determining different elements such as calcium, silicon and carbon.
  • Atomic Force Microscopy was used to compare the surface topography of hybrids fabricated in this study with a physical mixture of PMMA-bioglass.
  • the physical mixture (HO) possessed a non-uniform surface corroborating the appearance of separate phases of PMMA and bioglass.
  • the surface of the HI hybrid was homogenous and there was no evidence of phase separation both at low (5 ⁇ ) and high (1 ⁇ ) resolutions.
  • the surface roughness of HO was six- fold higher than that of the HI samples (see, Fig.18(c, f, j)).
  • the results of AFM also confirmed that the addition of MPMA as a coupling agent is efficient in creating a homogeneous hybrid from a polymer and bioglass and merging these two phases in the nanoscale range.
  • HI which promotes cells anchoring to the biomaterial surface and enhances cell spreading. Furthermore, for HI samples, cells were spread well on the surface and base lamellipodia were well-developed, thereby corroborating the adopted flattened morphology of the cell growth.
  • the degradation profile of the prepared hybrid was compared to that of the bioglass control samples. It was observed that presence of PMMA covalently bonded to the bioglass decreased the rate of degradation of the bioglass. As shown in Fig.24, linear lines were fitted to the degradation curves of the samples and it was observed that the degradation rate of the hybrid sample was 1.5 times less than the bioglass.
  • AFM analysis was used to compare the surface properties of hybrids fabricated in this study with physical mixture of PMMA-bioglass.
  • the roughness values were non-uniform for HO sample underlining the non- homogenous distribution of applied stress.
  • the surface of HI hybrid was smoother and the roughness dropped 5 fold compared to physical mixture.
  • Young's modulus of the HI that was prepared from 60:40 vol % PMFS:bioglass solutions were within the range suitable for osteoblast cells adhesion and proliferation. In vitro cell study was then conducted to assess the cell adhesion to this material.
  • the degradation of HI, bioglass and pure PMMA samples were measured within 100 days.
  • the data in Figure 27 shows that PMMA has no weight loss during this period; however, bioglass was rapidly degraded and completely dissolved after 83 days.
  • the mechanism of degradation of bioglass is well established in the literature.
  • Si0 2 is gradually dissolved and results in producing silicic acid, which is excreted from the body.
  • the degradation rate of Si0 2 matrices is a function of parameters such as composition of precursors, watenTEOS molar ratio, pH, and network connectivity of the silica matrix that is shown by Q n species.
  • Q n shows the number of Si-O-Si bonds around each silicon.
  • TEOS:water:HCl:CC l :8:0.01 :0.2 with 41% Q 3 and 53% Q 4 species that were calculated from 29 Si MR analysis.
  • the data in Figure 27 demonstrated that these characteristics led to the degradation rate of 0.92 % per day and 0.65 % per day for bioglass and HI samples, respectively.
  • the zero-order kinetics of bioglass degradation was in agreement with other studies.
  • the addition of PMMA to bioglass by covalent bonding (HI) resulted in impeding the degradation rate of bioglass for 1.5 fold and maintaining nearly 30 % of its original weight after 100 days of incubation. It was demonstrated that the mechanism of the new bone formation at the implant interface and the host tissue continues until there is ion exchange between the implant and body fluids.
  • the inorganic layer is well-distributed over the surface and it is more accessible for precipitation and formation of calcium-phosphate layer.
  • the organic phase covers the surface of samples and decreases the interface between inorganic phase and calcium-phosphate ions.
  • apatite on the surface of HI implies that the presence of inert PMMA polymer in the composition of 60:40 PMFS:bioglass had negligible impact on bioactivity.
  • Composite materials representing bulk mixing (HO) and MPMA-coupled hybrid (HI) of PMMA:bioglass were analyzed in vitro for their cell attachment and osteoblast growth properties. These materials were compared with their constituent materials, pure PMMA and bioglass. The bioglass was markedly brittle and prone to thermal and mechanical shock induced fracturing.
  • alkaline phosphatase (ALP) activity of samples was assessed.
  • Human pre-osteoblasts were grown upon HO, HI, PMMA and bioglass materials and treated with osteogenic differentiation media to induce expression of mature osteoblast markers.
  • Robust staining was seen on cells grown on tissue culture plastic as a positive control.
  • blue ALP+ cells were seen on the HI -treated hybrid samples but not on the HO samples or on the PMMA or bioglass alone.
  • mononuclear cells mainly macrophages and lymphocytes
  • macrophages and lymphocytes were embedded within a well vascularized granulation tissue.
  • the immunohistochemical detection of the macrophage-specific F4/80 antigen showed that these cells were mainly detectable as a monolayer at the material-tissue-interface and loosely distributed within the material-adherent granulation tissue.
  • TRIP tartrate-resistant acid phosphatase
  • multinucleated giant cells The occurrence of the latter was independent of fragment size. Overall, the implantation bed of this group was relatively extensive, which again can be attributed to the presence of the multinucleated giant cells, which are known to produce vascular endothelial growth factor (VEGF) among other substances responsible for material degradation. VEGF has negligible fragmentation during the observation period, which can be attributed to its higher mechanical stability. This material induced mainly mononuclear cells and a relatively low amount of multinucleated giant cells. Interestingly, the latter did not show a high TRAP-expression and it was observed that the implantation bed of HI was well vascularized.
  • VEGF vascular endothelial growth factor
  • PMMA-bioglass hybrids were produced by sol-gel method in the presence of chemical bonding that integrated the organic and inorganic components. This molecular level interaction addressed the issue of phase- separation that is commonly observed in preparation of physical mixtures of bioglass with a polymer.
  • the mechanical properties of the hybrid acquired at optimum composition of PMMA:MPMA was significantly improved compared with bioglass; the Young's modulus of hybrid was decreased 40 fold and its hardness was 16 fold higher than pure bioglass.
  • the chemical bonding of PMMA with bioglass resulted in prolonging the degradation of bioglass that may be favorable for bone regeneration and in situ drug release applications.
  • Precursors required for synthesis of PMMA-co-MPMA copolymer including MPMA, ⁇ , ⁇ '-Azoisobutyronitrile (AIBN) and ⁇ , ⁇ ' - dimethylformamide (DMF) were purchased from Sigma and used as received.
  • Methyl methacrylate (MMA) purchased from Sigma was used after distillation under reduced pressure.
  • Hydrochloric acid (HCl; Merck), tetraethyl orthosilicate (TEOS; Sigma), calcium chloride dihydrate (CaCl 2 .2H 2 0 (CC); Ajax Finechem Pty Ltd), tetrahydrofuran (THF; Merck) and deionized water were used for fabrication of the inorganic solution and the hybrid. 7.1.2 Preparation ofbioglass, PMMA-co-MPMA and pure PMMA solutions
  • TEOS was mixed with deionized water and HC1 and stirred for 30 minutes followed by addition of calcium chloride dihydrate.
  • a common calcium source for the preparation of sol-gel derived bioglasses is calcium nitrate tetrahydrate; however, in this study, calcium chloride was used to minimize the risk of toxicity resulted from nitrate by-product.
  • sol(B):sol(A) 60:40, then mechanically stirred for one hour to obtain a homogenous and well dispersed solution.
  • This composition was selected due to the fact that bioglass composition was shown to have no significant impact on the network characteristics and molecular integration of hybrids.
  • the product was then dried in vacuum oven at 40 °C for a period of 2 days. This temperature profile was developed for drying the samples to remove residues of solvents and maintain the monolith structure.
  • a thin film of polymer was obtained with a gel structure of bioglass that was subsequently grinded and well-mixed as powder. Afterwhich, the powder was dissolved in THF (10 wt%) and casted on a teflon container, which was then vacuum dried at 40 °C for 2 days.
  • Hybrid samples were prepared in the form of monoliths and uniaxial compression tests was performed in an unconfmed state with a 1000 N load cell by Instron (Model 5543). Dimensions of the samples were 6.61 ⁇ 0.05 mm diameter and 1.2 ⁇ 0.1 mm height. The samples were subjected to a loading and the Young's modulus was obtained as the tangent slope of the stress-strain curve between 0 % - 10 % strain level. The area under the compressive stress- strain curves was calculated for measuring the toughness of samples. Three samples were examined for each group for statistical analysis.
  • the degradation rate of samples was tested by measuring the change in sample weight over time under simulated physiological conditions. Three samples were kept in PBS at 37°C and at different time intervals they were removed from the degradation medium and rinsed three times with deionized water and dried prior to weighing. The measurement was continued for a period of 100 days.
  • the samples were characterized for their bioactivity by incubation in simulated body fluid (SBF) at 37 °C for the periods of 1, 3, 7, and 14 days. Samples were washed twice with deionized water to remove any residue of minerals absorbed on the surface. The dried monoliths were examined to determine the formation of apatite layer on the surface.
  • SBF simulated body fluid
  • HOB Human pre-osteoblasts
  • Complete osteoblast growth media (Invitrogen) was used to culture the HOB cells, which were incubated at 37 °C in the presence of 5% C0 2 and 95 % humidity. The media was refreshed every three days until the cells approached confluence.
  • the samples (12 mm diameter ⁇ 1 mm height) were placed into a well plate and kept in 70 % ethanol for one hour for sterilization, followed by rinsing with PBS three times. Samples were then exposed to UV light for 30 minutes and were washed with fresh medium at 37 °C overnight.
  • the substrates were placed into well-plates and seeded with cells at the density of 2 ⁇ 10 5 cells/ml and kept in culture for 7 days.
  • Cell morphology was examined at day four of culture by SEM (FE-SEM; Zeiss ULTRA plus) analysis after being fixed with glutaraldehyde according to previously published method. 64 Viability was measured at 1 day and 7 days post seeding. Cellular viability was assessed using the CellTitre 96 Aqueous One Solution Cell Proliferation Assay kit
  • samples were transferred to media supplemented with ascorbic acid (50 ⁇ g/ml), ⁇ -glycerophosphate (10 mM) and BMP-2 (200 ng/ml) after 48 hours cultured with cells.
  • ascorbic acid 50 ⁇ g/ml
  • ⁇ -glycerophosphate 10 mM
  • BMP-2 200 ng/ml
  • ALP is an osteogenic marker that is expressed by differentiating osteoblasts.
  • ALP staining was carried out on pure PMMA and bioglass, their physical mixture and hybrid at day 4. Samples were fixed with gluteraldehyde, washed with PBS and incubated in TRIS buffer (1 M, 9.4 pH) for 5 minutes. The scaffolds were then stained in Naphtol AS-BI phosphate (Sigma) as a substrate and Fast Blue (Sigma) as the stain. Scaffolds were washed with H 2 0 to remove excess staining before imaging. Cells alone on tissue culture plastic were used as a control. Images were captured using a Leica MZ6 microscope with a Qlmaging Micropublisher 5.0 camera. 7.3 In vivo study
  • the subcutaneous implantation of bone substitutes was applied according to a previously published operation procedure.
  • the animals were anesthetized with an intraperitoneal injection (10 ml of ketamine (50 mg ml "1 ) with 1.6 ml of 2% xylazine).
  • an intraperitoneal injection (10 ml of ketamine (50 mg ml "1 ) with 1.6 ml of 2% xylazine).
  • a subcutaneous pocket in the subscapular region was formed by means of a scalpel and surgical scissor.
  • the bone substitute materials were subsequently inserted under sterile conditions into the subcutaneous tissue pocket under the thin skin muscle of the subscapular region. Wound closure was performed by means of Prolene 6.0 suture material (Ethicon, Germany).
  • tissue preparation for all of the groups was performed according to a previously described method. Experimental animals were sacrificed by an overdose of Ketamine and Xylazin at day 10 after implantation. After 10 days the bone substitute materials were resected together with the surrounding peri- implant tissue. Tissue fixation was carried out by means of 4% formalin for 24 hours. For further histological workup and (immuno-) histochemical staining, the tissue of the implant site was cut into three segments of identical dimensions containing the left margin, the center and the right margin of the biomaterial. Paraffin embedding was performed after dehydration of the biopsies in a series of increasing alcohol concentrations followed by xylol incubation. Six ongoing 3-5- ⁇ thick sections were made from the central segment of each animal by means of a rotation microtome.
  • the material-tissue interaction was visualized by means of previously published histochemical and immunohistochemical staining methods.
  • the first three slides sections were stained with haematoxylin and eosin (H&E), Movat ' s Pentachrome and Azan, respectively.
  • the fourth slide was used to identify osteoclast-like cells by tartrate-resistant acid phosphatase (TRAP) staining according to previously described methods.
  • the fifth slide was used for immunochemical staining with a pan-macrophage marker (F4/80 antibody, rat monoklonal, Clon BM8, eBioScience, USA) in combination with peroxidase and diaminobenzidine (En Vision Detection System, Peroxidase/DAB, rabbit/mouse, K5007; Dako Cytomation, Hamburg, Germany).
  • the sixth slide served as a control of the staining method in absence of the F4/80 antibody. All of the other chemicals were purchased from Sigma- Aldrich and used without further purification.
  • the histopathological evaluation was performed by two independent investigators (SG and MB) by means of a conventional diagnostic microscope (Nikon Eclipse 80i, Tokyo, Japan). The description and the outcome of the cell- and tissue-biomaterial interactions were evaluated by examination of the total implantation bed and its peri-implant tissue as previously described.
  • PMMA-bioglass composites were fabricated through the nano-scale interaction between the organic and inorganic components using silane coupling agent at optimized condition of
  • porous scaffolds were produced with 67% porosity and interconnective pores in the range of 100-300 ⁇ by using nontoxic SB at optimum concentration.
  • the alizarin red S staining showed the differentiation of osteoblast cells on fabricated PMMA-bioglass hybrid underpinning its favorable characteristics for bone replacement applications.
  • the procedure of preparing the hybrids disussed above included dissolving MPMA-functionalized polymer in tetrahydrofuran (THF), which was subsequently mixed with bioglass sol. Fabricated transparent hybrid solution was gelled within 5 hours at room temperature and then dried for a period of 2 weeks for hardening and forming a monolithic crack-free structure. Removing the impurities such as TFIF and acid was also conducted during drying stage.
  • THF tetrahydrofuran
  • sol-gel method was used for the fabrication of porous scaffolds from hybrid class of material.
  • the pore formation in sol-gel derived hybrid solution was usually conducted by adding a surfactant such as Triton X- 100 to produce bubbles when agitating followed by the addition of hydrofluoric acid (FIF) as the catalyst for condensation of silica and fixative for bubbles.
  • FIF hydrofluoric acid
  • Fast gelation time of hybrid solution ( ⁇ 10 minutes) is an important parameter, which determines the feasibility of this technique by fixation of bubbles inside the structure.
  • Mouse clonal pre-osteoblast cells (MC3T3) were used to assess the cell interaction with samples.
  • Alpha Minimum Essential Medium (a-MEM) supplemented with 10% fetal bovine serum (FBS), 1% L-glutamine and 1% penicillin/streptomycin was used to culture the cells, which were incubated at 37°C in the presence of 5% C0 2 and 95 % humidity. The media was refreshed every three days until the cells approached confluence.
  • MC3T3 cells were detached from flask by trypsin treatment and collected for cell seeding.
  • the monolith samples (6 mm diameter x 3 mm height) and porous scaffolds (12 mm diameter x 2 mm height) were placed into well plates and kept in 70 % ethanol under UV light for one hour for sterilization followed by rinsing with PBS triple.
  • the samples were placed into well tissue-culturing polystyrene plates (TCPS) and cells with the density of 1 ⁇ 10 5 cells/ml were seeded onto the monoliths and porous scaffolds, respectively.
  • TCPS was used as control in this study.
  • the samples were subsequently kept in a C0 2 incubator at 37 °C for further analyses.
  • samples were transferred to media supplemented with ascorbic acid (50 ⁇ / ⁇ 1), ⁇ -glycerophosphate (10 mM) and BMP-2 (200 ng/ml) after 7 days cultured with cells. The time of transferring the samples to this media was considered as day 0 for viability with osteogenic media.
  • Viability was measured at day 7 after seeding cells and the addition of osteogenic media. Cellular viability was assessed using the CellTitre 96
  • Aqueous One Solution Cell Proliferation Assay kit (Promega) according to the manufacturer's instructions. Briefly, scaffolds were incubated with the viability solution for 30 min at 37 °C and read using a spectrophotometer at 595 nm. Cell growth and proliferation was measured by viability assay on cultured samples at day 1 and day 7 after cell seeding. Furthermore, the viability of cells after the addition of osteogenic media was also assessed after 7 days. All samples were assayed in triplicate for statistical analysis.
  • the formation of calcium phosphate was determined by using Alizarin Red S staining after 28 days of cell culture (21 days after the addition of OM). Alizarin Red S-calcium complex is formed in this assay in a chelation process, and the end product is birefringent. Scaffolds were fixed in 4%
  • PMMA is soluble in solvents such as THF, acetic acid, l-Methyl-2- pyrrolidone (NMP) and ethanol at specific conditions.
  • solvents such as THF, acetic acid, l-Methyl-2- pyrrolidone (NMP) and ethanol at specific conditions.
  • ethanol was used as an alternative to THF for the fabrication of PMMA-bioglass hybrids to assess the effect of solvent due to its lower boiling temperature ⁇ i.e. 78 °C) compared with NMP (with boiling temperature of 202 °C) and acetic acid.
  • ethanol has been shown to be safe for biomedical formulations.
  • SAIB sucrose acetate isobutyrate
  • PMMA-co-MPMA copolymer was soluble in ethanol at above 70 °C.
  • the addition of bioglass solution inhibited the precipitation of this copolymer in ethanol at below 70 °C. This effect was due to the rapid covalent bonding between this copolymer and bioglass sol.
  • This temperature was, therefore, used for the formation of chemical bonding between silanol groups of MPMA and bioglass.
  • the hybrid solution was then rapidly cooled down to a predetermined temperature for further processing and characterization.
  • the optimized PMMA-bioglass hybrid with 0.1 molar ratio MPMA:MMA and 60:40 vol% polymenbioglass solutions was used in this study.
  • the effect of solvent on gelation time was compared when the hybrids (HI and Hl E toH) were prepared at room temperature (25 °C).
  • the gelation time of samples was measured by test tube tilting method. As shown in Figure 37, the gelation time of Hl E tOH hybrid was 120 ⁇ 7 minutes that was significantly lower than HI sample that required 300 ⁇ 12 minutes (at 25 °C) to gel.
  • the gelation time of HI and Hl E tOH was reduced from 300 ⁇ 12 minutes to 90 ⁇ 7 minutes and from 120 ⁇ 7 minutes to 37 ⁇ 2 minutes, respectively.
  • the gelation time of Hl Et o H hybrid was reduced from 120 ⁇ 7 minutes to 7 ⁇ 0.03 minutes. It is important to note that it was not practical to fabricate hybrid in THF (HI) at above 55 °C due to the rapid solvent evaporation and precipitation of copolymer in sol-gel system.
  • SB Sodium bicarbonate
  • SB can be used as an alternative to HF that is commonly used in sol-gel method for accelerating the condensation of silanol groups and reducing the gelation time. SB has no toxicity compared to HF that is counted as a corrosive chemical.
  • Hl E tOH was produced at optimized condition with fastest gelation time.
  • the network factors (Q n ) were measured to determine the impact of gelation time on the structure of final product.
  • Q n shows a silicon bonds to n other silicons via "bridging" oxygen atoms.
  • the Q n species and network connectivity model (Nc) of samples (HI and Hl E toH) were listed in Table 10.
  • Nc which ranges between zero to 4 describes the characteristics of silica network and is defined as the average number of bridging oxygen atoms per silicon 8 . It was shown that the acceleration of gelation time resulted in increasing the number of Q 4 species and Nc value, underpinning constructing a more condensed structure.
  • the Nc of Hl E tOH was enhanced 15 % compared with HI .
  • the desired mechanical properties for bone replacement are to possess low Young's modulus while maintaining the high ultimate stress. It was demonstrated in our previous study that the integration of PMMA into the structure of bioglass resulted in 40-fold decrease in the Young's modulus and this value was decreased from 229 MPa in pure bioglass to 6 MPa in HI samples. This effect was efficient in addressing the issue of brittleness of bioglass, however, the ultimate stress of HI sample was still not sufficient for load-bearing bone replacement applications. In this study, the effect of fast network formation and more condensed structure on the mechanical properties of samples was investigated. As shown in Figure 39, reducing the gelation time in Hl E toH had negligible effect on the Young's modulus of samples compared with HI .
  • Hl E tOH was 5.6-fold higher than HI and approached 5.7 ⁇ 1.16 MPa. This value was in the range of the ultimate stress of cancellous bone (i.e. 5-10 MPa). However, it was lower than the ultimate stress of cortical bone (i.e. 100-230 MPa) and sintered bioglass at 800 °C (70S30C) (i.e. 32-89 MPa).
  • Hl E toH was 1.6-fold greater than HI due to the fast gelation time and more condensed silica structure (higher Nc value). Higher toughness value of Hl E toH is beneficial as the sample is more resistible to the applied load.
  • Hybrid solution was injected through a 21G needle and the injectability profile at room temperature was shown in Figure 41. Mixing and loading the syringe were carried out in 2 minutes. Subsequently the syringe was exposed to the load (N) and the stress versus time was recorded. Solution was steadily extruded through the needle and no increase was detected in the load until 60 seconds. After which, a sudden increase was observed in the required load, due to gelation of hybrid solution.
  • This solution can be contemplated as an alternative for injectable polymers mixed with hydroxyapatite or bioglass particles because this system reduced the risk associated with needle blockage due to the presence of bioglass particles.
  • the hybrid also provides a more homogenous mixture of polymer and bioglass.
  • the degradation property of bone implants is one of the most challenging issues in the application of composite materials. Inhomogeneous degradation of glassy/ceramic and polymer compounds results uneven stress distribution or load transmission on the composite materials that leads to adverse effects such as loosening of the implant. It was demonstrated that the formation of nano-scale interaction between organic and inorganic components addressed the issue of separate degradation and hybrids possessed superior degradation properties compared with physical mixtures and conventional composites.
  • the degradation mechanism of bioglass is well established in the literature. Briefly, The Si-O-Si network of bioglass gradually disrupts producing silicic acid, which is excreted from the body. This phenomena is a function of parameters such as composition of precursors, watenTEOS molar ratio, pH, and Nc. For example, low Nc value results in less compact network, hence faster degradation rate. Furthermore, it was demonstrated that the release of Si ions to body fluids is beneficial for the mechanism of new bone formation at the site of implant .
  • Hl Et oH can be an appropriate candidate for long term bone implant application due to the high mechanical properties, long- term ion exchange with surrounding tissue and the ability to be shaped as required.
  • SB The low concentrations of SB (e.g. 0.4 wt%) was used to form firm monolithic structures. However, higher concentrations (e.g. 10 wt%) was used for the fabrication of porous scaffolds. SB possessed dual effect: a) formation of bubbles due to the decomposition of SB at 70 °C into C0 2 , and b) rapid gelation (15 seconds) due to increase in the pH of hybrid solution, which fixed the bubbles. During gelation the pores were also ruptured leading to interconnectivity in the structure. As shown in Figure 43(a), the pores were irregular and their size ranged between 100-300 ⁇ , which is in the range suitable for osteoblast cells proliferation.
  • Porous scaffolds fabricated from Hl E toH formulation were also examined.
  • Porous scaffolds (PS) similar to all other porous structures, possessed lower value of Young's modulus and compressive strength compared to monoliths.
  • the Young's modulus of porous scaffolds was 0.087 ⁇ 0.006 MPa and their toughness value was 4.57 ⁇ 1.14 Nm "2 .
  • the stress- strain diagram of these samples was very steady and the samples were as resistible as the monoliths to the level of strain. As shown in Figure 44, the ultimate strain of porous samples was 0.26 mm/mm, which showed negligible difference with the value for monolithic structures ⁇ i.e. 0.32 mm/mm).
  • porous scaffolds with high pore interconnectivity were fabricated during sol-gel technique by using SB as the gas foaming and silica condensation catalyst.
  • the differentiation of osteoblast cells on the hybrids was confirmed and opened an avenue for the application of fabricated hybrid for replacing damaged bone tissue.

Landscapes

  • Health & Medical Sciences (AREA)
  • Chemical & Material Sciences (AREA)
  • Life Sciences & Earth Sciences (AREA)
  • Veterinary Medicine (AREA)
  • Public Health (AREA)
  • General Health & Medical Sciences (AREA)
  • Animal Behavior & Ethology (AREA)
  • Epidemiology (AREA)
  • Oral & Maxillofacial Surgery (AREA)
  • Transplantation (AREA)
  • Medicinal Chemistry (AREA)
  • Dermatology (AREA)
  • Dispersion Chemistry (AREA)
  • Inorganic Chemistry (AREA)
  • Engineering & Computer Science (AREA)
  • Composite Materials (AREA)
  • Materials Engineering (AREA)
  • Heart & Thoracic Surgery (AREA)
  • Surgery (AREA)
  • Vascular Medicine (AREA)
  • Materials For Medical Uses (AREA)

Abstract

La présente invention concerne un matériau composite comprenant un polymère-bioverre couplé de façon covalente, le polymère étant fonctionnalisé avec un agent de couplage ou comprenant un agent de couplage de comonomère adapté pour coupler de façon covalente le polymère et le bioverre. L'invention concerne en outre un procédé de production d'un matériau composite comprenant les étapes de : préparation d'un sol inorganique de précurseurs de bioverre; préparation d'un sol organique de polymère; combinaison du sol inorganique et dudit sol organique dans des conditions permettant de produire un matériau composite polymère-bioverre couplé de façon covalente. Le procédé concerne en outre la formation d'un matériau composite in situ ou in vivo, comprenant les étapes d'administration à un sujet ayant un tel besoin d'un sol de polymère et d'un sol de bioverre à un emplacement prédéterminé dans le corps du sujet et catalyse in situ de la formation d'un matériau composite bioverre-polymère hybride. Le procédé concerne en outre une formulation pour préparer un matériau composite, et un kit comprenant dans des récipients séparés : un sol de bioverre, un sol de copolymère et un catalyseur. L'invention concerne en outre l'utilisation d'un composite hybride en tant que matériau de remplissage osseux, sous la forme d'une formulation injectable, ou sous la forme d'un échafaudage poreux pour des applications biomédicales. La présente invention concerne un procédé pour améliorer la stabilité à long terme d'un dispositif médical implantable, et un procédé pour la régénération ou le resurfaçage de tissu, comprenant les étapes de : mise en contact dudit tissu avec le composite hybride de l'invention.
PCT/AU2014/000126 2013-02-14 2014-02-14 Matériau biocompatible et utilisations de celui-ci WO2014124496A1 (fr)

Applications Claiming Priority (2)

Application Number Priority Date Filing Date Title
AU2013900475 2013-02-14
AU2013900475A AU2013900475A0 (en) 2013-02-14 Improved PMMA-bioglass hybrid networks

Publications (1)

Publication Number Publication Date
WO2014124496A1 true WO2014124496A1 (fr) 2014-08-21

Family

ID=51353430

Family Applications (1)

Application Number Title Priority Date Filing Date
PCT/AU2014/000126 WO2014124496A1 (fr) 2013-02-14 2014-02-14 Matériau biocompatible et utilisations de celui-ci

Country Status (1)

Country Link
WO (1) WO2014124496A1 (fr)

Cited By (14)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
CN107213528A (zh) * 2017-05-31 2017-09-29 苏州蔻美新材料有限公司 一种可降解骨植入物的制备方法
CN107789096A (zh) * 2016-09-07 2018-03-13 艾尔生技有限公司 生医支架及其制造方法
US10238507B2 (en) 2015-01-12 2019-03-26 Surgentec, Llc Bone graft delivery system and method for using same
CN110177636A (zh) * 2016-11-15 2019-08-27 霍加纳斯股份有限公司 用于增材制造法的原料、使用其的增材制造法和由其获得的制品
WO2019232589A1 (fr) * 2018-06-07 2019-12-12 The University Of Sydney Implant biodégradable pour des applications orthopédiques, ayant des caractéristiques spécialement ajustables
US10687828B2 (en) 2018-04-13 2020-06-23 Surgentec, Llc Bone graft delivery system and method for using same
CN111458416A (zh) * 2019-01-18 2020-07-28 辽宁远大诺康生物制药有限公司 聚对二氧环己酮中杂质的检测方法
CN111905156A (zh) * 2019-05-10 2020-11-10 华东理工大学 一种高强度可吸收活性复合材料及其制备方法
CN112204093A (zh) * 2018-06-25 2021-01-08 国家航空航天技术研究所 用于降解检测的同位素标记材料
US11116647B2 (en) 2018-04-13 2021-09-14 Surgentec, Llc Bone graft delivery system and method for using same
CN113698094A (zh) * 2021-08-31 2021-11-26 北京航空航天大学 一种高强高韧高透光率复合材料的制备方法
CN114460159A (zh) * 2022-02-17 2022-05-10 河南中医药大学 基于photo-ATRP信号放大策略的ALP活性检测试剂盒及其使用方法
CN117303736A (zh) * 2023-09-18 2023-12-29 哈尔滨理工大学 一种含有机硅组份的生物活性玻璃及其制备方法
CN118078644A (zh) * 2024-04-23 2024-05-28 北京大学口腔医学院 一种牙科复合树脂及其制备方法和应用

Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4652459A (en) * 1982-11-10 1987-03-24 Achim Engelhardt Implants, and process for the production thereof
US6399693B1 (en) * 1997-10-23 2002-06-04 Univ. Of Florida Research Foundation Bioactive composites comprising silane functionalized polyaryl polymers
US20030113686A1 (en) * 2001-10-24 2003-06-19 Weitao Jia Root canal filling material
EP2243500B1 (fr) * 2009-04-23 2012-01-04 Vivoxid Oy Composite biocompatible et son utilisation
CN102886069A (zh) * 2012-09-24 2013-01-23 华南理工大学 溶胶凝胶生物玻璃-高分子杂化材料的制备方法

Patent Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4652459A (en) * 1982-11-10 1987-03-24 Achim Engelhardt Implants, and process for the production thereof
US6399693B1 (en) * 1997-10-23 2002-06-04 Univ. Of Florida Research Foundation Bioactive composites comprising silane functionalized polyaryl polymers
US20030113686A1 (en) * 2001-10-24 2003-06-19 Weitao Jia Root canal filling material
EP2243500B1 (fr) * 2009-04-23 2012-01-04 Vivoxid Oy Composite biocompatible et son utilisation
CN102886069A (zh) * 2012-09-24 2013-01-23 华南理工大学 溶胶凝胶生物玻璃-高分子杂化材料的制备方法

Non-Patent Citations (2)

* Cited by examiner, † Cited by third party
Title
BAVARIAN, R. ET AL.: "Improving the Bioactivity of Bioglass / (PMMA-co-MPMA) Organic / Inorganic Hybrid", ENGINEERING IN MEDICINE AND BIOLOGY SOCIETY, EMBC, 2011 ANNUAL INTERNATIONAL CONFERENCE OF THE IEEE., 30 August 2011 (2011-08-30), Retrieved from the Internet <URL:http//ieexploreeeeorg//articleDetails.lsp?ber=6090601> [retrieved on 20140331] *
RAVARIAN, R. ET AL.: "Molecular interactions in coupled PMMA-bioglass hybrid networks", JOURNAL OF MATERIALS CHEMISTRY B, vol. 1, 2013, pages 1835 - 1845, Retrieved from the Internet <URL:http/publsrsc.org/en/content/articiclanding/2013/tb/c2b00251e#ldivAbstract> [retrieved on 20140331] *

Cited By (18)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US10238507B2 (en) 2015-01-12 2019-03-26 Surgentec, Llc Bone graft delivery system and method for using same
US11116646B2 (en) 2015-01-12 2021-09-14 Surgentec, Llc Bone graft delivery system and method for using same
CN107789096A (zh) * 2016-09-07 2018-03-13 艾尔生技有限公司 生医支架及其制造方法
CN107789096B (zh) * 2016-09-07 2021-10-26 艾尔生技有限公司 生医支架的制造方法
CN110177636A (zh) * 2016-11-15 2019-08-27 霍加纳斯股份有限公司 用于增材制造法的原料、使用其的增材制造法和由其获得的制品
CN107213528A (zh) * 2017-05-31 2017-09-29 苏州蔻美新材料有限公司 一种可降解骨植入物的制备方法
US11116647B2 (en) 2018-04-13 2021-09-14 Surgentec, Llc Bone graft delivery system and method for using same
US10687828B2 (en) 2018-04-13 2020-06-23 Surgentec, Llc Bone graft delivery system and method for using same
WO2019232589A1 (fr) * 2018-06-07 2019-12-12 The University Of Sydney Implant biodégradable pour des applications orthopédiques, ayant des caractéristiques spécialement ajustables
CN112204093A (zh) * 2018-06-25 2021-01-08 国家航空航天技术研究所 用于降解检测的同位素标记材料
CN111458416A (zh) * 2019-01-18 2020-07-28 辽宁远大诺康生物制药有限公司 聚对二氧环己酮中杂质的检测方法
CN111458416B (zh) * 2019-01-18 2022-08-09 沈阳恒生医用科技有限公司 聚对二氧环己酮中杂质的检测方法
CN111905156A (zh) * 2019-05-10 2020-11-10 华东理工大学 一种高强度可吸收活性复合材料及其制备方法
CN113698094A (zh) * 2021-08-31 2021-11-26 北京航空航天大学 一种高强高韧高透光率复合材料的制备方法
CN114460159A (zh) * 2022-02-17 2022-05-10 河南中医药大学 基于photo-ATRP信号放大策略的ALP活性检测试剂盒及其使用方法
CN114460159B (zh) * 2022-02-17 2023-11-03 河南中医药大学 基于photo-ATRP信号放大策略的ALP活性检测试剂盒及其使用方法
CN117303736A (zh) * 2023-09-18 2023-12-29 哈尔滨理工大学 一种含有机硅组份的生物活性玻璃及其制备方法
CN118078644A (zh) * 2024-04-23 2024-05-28 北京大学口腔医学院 一种牙科复合树脂及其制备方法和应用

Similar Documents

Publication Publication Date Title
WO2014124496A1 (fr) Matériau biocompatible et utilisations de celui-ci
Dorozhkin Calcium orthophosphate-based biocomposites and hybrid biomaterials
Dorozhkin Biocomposites and hybrid biomaterials based on calcium orthophosphates
Cui et al. The nanocomposite scaffold of poly (lactide-co-glycolide) and hydroxyapatite surface-grafted with L-lactic acid oligomer for bone repair
Kerativitayanan et al. Elastomeric and mechanically stiff nanocomposites from poly (glycerol sebacate) and bioactive nanosilicates
Vergnol et al. In vitro and in vivo evaluation of a polylactic acid‐bioactive glass composite for bone fixation devices
Liu et al. Surface modification of bioactive glass nanoparticles and the mechanical and biological properties of poly (L-lactide) composites
Hasan et al. In vitro and in vivo evaluation of bioglass microspheres incorporated brushite cement for bone regeneration
Liu et al. Novel injectable calcium phosphate/chitosan composites for bone substitute materials
Shaltooki et al. Fabrication of chitosan-coated porous polycaprolactone/strontium-substituted bioactive glass nanocomposite scaffold for bone tissue engineering
Davison et al. Degradation of biomaterials
Niiranen et al. In vitro and in vivo behavior of self‐reinforced bioabsorbable polymer and self‐reinforced bioabsorbable polymer/bioactive glass composites
Kucko et al. Tough and osteocompatible calcium phosphate cements reinforced with poly (vinyl alcohol) fibers
Lima et al. Injectable bone substitute based on chitosan with polyethylene glycol polymeric solution and biphasic calcium phosphate microspheres
US8034850B2 (en) Self-setting absorbable composites and applications thereof
Victor et al. Bioactive, mechanically favorable, and biodegradable copolymer nanocomposites for orthopedic applications
Shahbazi et al. Introducing an attractive method for total biomimetic creation of a synthetic biodegradable bioactive bone scaffold based on statistical experimental design
Victor et al. Polymer ceramic composite materials for orthopedic applications—relevance and need for mechanical match and bone regeneration
Lagarrigue et al. Poly (d, l-lactide)-grafted bioactive glass nanoparticles: from nanobricks to freeze-cast scaffolds for bone substitution
Ohtsuki et al. Design of novel bioactive materials through organic modification of calcium silicate
Boesel et al. The in vitro bioactivity of two novel hydrophilic, partially degradable bone cements
Nerantzaki et al. A biomimetic approach for enhancing adhesion and osteogenic differentiation of adipose-derived stem cells on poly (butylene succinate) composites with bioactive ceramics and glasses
Szczepanska et al. Influence of ceramic phosphate powders on the physicochemical and biological properties of Poly (L-lactide)
Wan et al. Enhanced degradability of the apatite-based calcium phosphate cement incorporated with amorphous MgZnCa alloy
Ahmadi et al. Evaluation of hMSCs response to sodium alginate/bioactive glass composite paste: effect of CaO/P2O5, sodium alginate concentration and P/L ratios

Legal Events

Date Code Title Description
121 Ep: the epo has been informed by wipo that ep was designated in this application

Ref document number: 14751910

Country of ref document: EP

Kind code of ref document: A1

NENP Non-entry into the national phase

Ref country code: DE

122 Ep: pct application non-entry in european phase

Ref document number: 14751910

Country of ref document: EP

Kind code of ref document: A1