DELTA-LACTONE

Field of the Invention

The present invention concerns the field of synthetic bone cements useful as filler in oral and maxillofacial as well as orthopedic surgery. More in particular the present invention relates to a new injectable calcium phosphate cement comprising a so-called pore forming agent. In addition, the present invention provides the uses of the new bone filler cement in oral and maxillofacial as well as orthopedic intervention.

Background of the Invention

Injectable calcium phosphate cements (CPCs) are frequently applied as synthetic bone fillers in oral and maxillofacial as well as orthopedic surgery owing to their excellent clinical handling properties and superior bone response. Calcium phosphate minerals, specifically hydroxyapatite, constitute 70 % of bone tissue. Synthetic hydroxyapatites have shown to stimulate osteoblast adhesion and migration in several in vitro studies. Complex 3D hydroxyapatitic scaffolds can be fabricated by different techniques, however they are not easily applicable in locations with difficult accessibility such as maxillofacial bone defects. In 1986 calcium phosphate cement (CPC) was developed by L. C. Chow and W. E. Brown. Calcium phosphate powder when mixed with an aqueous solution generates a moldable paste that after injection can be shaped to fit the bone defect contours whereafter it will set to form a 3D scaffold. Apatitic calcium phosphate cements are preferred in both science and industry owing to the resemblance with the mineral phase of bone and teeth.

Apatitic cements however exhibit very slow degradation rates that hamper bone ingrowth and thus regeneration of bone tissue. Compared to the gold standard in bone- regenerating therapies (i.e. bone grafting using patient's own bone that is actively involved in bone remodeling), insufficient degradability of CPCs is a major drawback which strongly reduces its attractiveness for reconstructive surgery.

In order to solve this problem, over the past decades, research has been focusing on creating macroporosity to accelerate the degradation rate of the calcium phosphate ceramics in general and allow osteolytic cells to infiltrate the pores. It has been proven in various in vivo implantation studies that macroporosity within CPCs is accompanied by bone ingrowth into the (interconnected) pores while the degradation of the CPC artificial implants is accelerated. In attempts to enhance CPC replacement by newly formed bone, porogens of different nature have been included in CPC, such as calcium sulphate, biodegradable polymers, foaming agents, water soluble additives etc. The majority of these strategies has only been applied successfully to pre-made calcium phosphate ceramics. Traditional polymeric porogens, which create porosity upon removal by means of firing or dissolution using organic solvents are not suitable for injectable CPCs. Porosity generation by means of salt leaching is also incompatible with CPCs since the liquid component of CPCs dissolves these inorganic porogens prior to setting of the cement in the bone defect site.

Some porogens, such as poly-trimethylene-carbonate and gelatin, undergo enzymatic digestion that leads to slow surface erosion and degradation in situ. Gelatin is well known for its excellent biocompatibility and therefore it is being used in various medical areas, but its use as CPC porogen was unsuccessful due to slow enzymatic degradation .

The most successful approach for introducing macroporosity in situ, so far, has been the introduction of degradable polymeric microspheres based on polylactic-co-glycolic acid (PLGA). The lactic and glycolic acidic monomers produced upon PLGA degradation are known to enhance calcium phosphate cement dissolution both in vivo and in vitro. When porogens degrade at a faster rate than the CPC, a porous scaffold is obtained which allows fluid flow and new tissue ingrowth within the 3D scaffolds enhancing its degradation rate.

Despite the fact that, CPC in combination with PLGA microspheres have shown enhanced degradation and good new bone formation through the scaffold, the process is still slow, taking between several months up to years depending on the exact chemical nature of the polymer.

Furthermore, the development of an (industrial) process of producing PLGA microspheres of appropriate size to generate pores that are sufficient for tissue ingrowth while not compromising the architecture of the material, at present, is proving cumbersome. This increases the complexity of the CPC formulation and reduces the feasibility of large-scale commercial application of CPCs containing processed porogens.

Therefore, a great demand exists for improved CPCs that can be injected and wherein macroporosity starts to develop within days rather than weeks or months. Ideally, porosity should be created in situ within CPCs directly after injection/setting of the cement to maximize the amount and rate of bone ingrowth, without negatively affecting the setting and biological properties of the cement. Furthermore, an ideal porogen has to be biocompatible, cost-effective, off-the-shelf available and exhibit good handling properties to be used by the surgeons in the operation room. It is the objective of the present invention to provide a CPC that constitutes a significant improvement with respect to one or more of these characteristics.

Summary of the Invention

The present inventors have surprisingly found that this objective can be realized by incorporating, in a calcium phosphate cement, a pore forming agent based on glucono-delta- lactone (GDL).

GDL, the cyclic ester of gluconic acid, undergoes hydrolysis in the physiologic environment leading to the production of gluconic acid which is soluble in water. This process results in voids at the space initially taken in by the solid GDL material. Additionally, acid generation accelerates the degradation of the CPC matrix creating additional macroprosity. In this process, the rate of GDL hydrolysis is of major importance. When hydrolysis would occur at a very high rate, setting of the CPC may be compromised, if it would occur at a very slow rate the CPC degradation rate would become insufficient, as is the case with, for instance, PLGA.

The present inventors found that incorporation of GDL based pore forming agent in conventional CPC leads to acidification of the surrounding CPC and results in a macroporous CPC structure within days. GDL displays a dissolution rate that is much faster than PLGA. At the same time, the gluconic acid is not liberated at a rate so high that it interferes with the setting reaction of the CPC.

Furthermore, cohesion tests have shown that GDL-CPC maintains its injectability properties when injected directly in aqueous media. As a matter of fact, the injectability of GDL/CPC composites was improved compared to PLGA/CPC composites in terms of injection force and subsequent setting.

As indicated above, pore formation by GDL does not solely rely on the formation of

"latent pores" in the setting cement leaving pores after it diffuses, dissolves, or degrades. In accordance with the present invention, formation of the desired macroporosity, at least in part, is based on CPC degradation by the acid liberated from GDL. This means that the size and shape of the of pores formation is not confined to the space initially reserved by the pore forming agent so that it is not necessary to provide the GDL in the form of microspheres with a predefined size and shape to obtain the desired macroporosity. As a matter of fact, GDL products in a form suitable for use in accordance with the invention are already commercially available. The appending example shows the effects of GDL as pore forming agent in CPC' s in vitro and in vivo, confirming that this novel material combination is superior as bone filler material to other hydrolytically degradable composites (CPC-PLGA) and enzymatically degradable composites (CPC-gelatin).

GDL-CPC is highly biocompatible. Its end-product gluconic acid can be naturally eliminated by the body. It is a metabolite that occurs after oxidation of glucose. No major adverse reactions are observed when using GDL-CPC in vivo in rabbits. At this moment there is no other alternative material available with similar properties as the combination CPC-GDL in terms of biodegradability and biocompatibility.

Hence, a first aspect of the invention concerns a bone cement powder useful as bone cement comprising a combination of one or more calcium phosphate compounds and a particulate glucono-delta-lactone material.

A second apect of the invention concerns a bone cement comprising a combination of one or more calcium phosphate compounds and a particulate glucono-delta-lactone material, and an aqueous liquid phase.

A third aspect of the invention concerns the use of a bone cement or bone cement powder, comprising a combination of one or more calcium phosphate compounds and a particulate glucono-delta-lactone material, in a method of oral, maxillofacial or orthopedic surgery in a subj ect in need thereof

A fourths aspect of the invention concerns a method of oral, maxillofacial or orthopedic surgery in a subject, said method comprising:

i) providing a bone cement comprising a combination of one or more calcium phosphate compounds and a particulate glucono-delta-lactone material, and an aqueous liquid phase;

ii) applying the bone cement to a defect or cavity in a tissue of said subject so as to fill said defect or cavity; and

iii) allowing the bone cement to set.

A fifth aspect of the invention concerns a kit of parts comprising at least two reservoirs comprising distinct compositions which can be used in preparing a bone cement of the invention, e.g. a first reservoir holding a bone cement powder according to the invention and a second reservoir holding a suitable liquid phase.

These and other aspects of the invention are described in more detail in the following sections. Detailed description of the Invention

As used herein, the term "bone cement powder" refers to a mixture that is the precursor of cement. The mixture can be in a dry powder or granular form. Upon mixing with a suitable liquid, the mixture forms a plastic paste. The paste undergoes a chemical reaction and/or a crystal rearrangement and hardens with time into a set solid cement as a result of the hydration reaction. The liquid initiator can be any physiologically acceptable aqueous system, e. g. water, an aqueous buffer or an aqueous solution. Preferably, the bone cement powder of this invention is a mixture comprising a calcium phosphate component and the GDL based pore forming agent.

Self-setting calcium phosphate cement scan be categorized as either brushite-forming or apatite-forming cements. In a preferred embodiment of this invention the calcium phosphate component is apatatite forming. In an even more preferred embodiment of the invention, the calcium phosphate component is hydroxy-apatite forming.

In accordance with an embodiment of the invention, the calcium phosphate component is based on one or more calcium phosphate compounds with or without additional calcium salts. The calcium phosphate compounds useful in the invention include hydroxyapatite (HA), monocalcium phosphate (MCP), amorphous calcium phosphate (ACP), monocalcium phosphate monohydrate (MCPM), dicalcium phosphate (DCP), dicalcium phosphate dihydrate (DCPD), dicalcium phosphate anhydrous (DCPA), precipitated or calcium-deficient apatite (CD A), alpha- or beta-tricalcium phosphate (a-TCP, β-TCP), tetracalcium phosphate (TTCP), octacalcium phosphate, BIOGLASS™, fluorapatite, chlorapatite, magnesium- substituted tricalcium phosphate, carbonate hydroxyapatite, other substituted forms of hydroxyapatite (e.g., hydroxyapatite derived from bone may be substituted with other ions such as fluoride, chloride, magnesium, sodium, potassium, etc.), and combinations and derivatives thereof. Easily resorbable calcium phosphate compounds are preferred.

A calcium phosphate component selected from the group consisting of, a-TCP, TTCP, ACP, MCPM, DCPA and mixtures thereof, is particularly preferred. A calcium phosphate component comprising a-TCP or TTCP is more preferred. A calcium phosphate component comprising a-TCP is most preferred.

a-TCP has the formula a-Ca3(P0 ₄ )2. a-TCP is easily transformed into calcium- deficient hydroxyapatite (CDA) in aqueous solution. This property is used to form apatitic CPCs. An "apatitic" calcium phosphate cement crystallises in the hexagonal system having the formula Ca _5x (P0 ₄ ) ₃ x,(OH,Cl,F) _x or (C0 ₃ )o _.5x . A calcium phosphate component may consist of a single calcium phosphate compound or a combination of two, three or more calcium phosphate compounds.

In an exemplary embodiment of the invention the calcium phosphate component consists essentially of a-TCP. In an exemplary embodiment of the invention the calcium phosphate component consists essentially of TTCP. Suitable combinations of calcium phosphate compounds include a-TCP, DCPA and precipitated hydroxyapatite; a-TCP and ACP; β-TCP and MCPM; TCP, DCPD and calcium carbonate CC; MCP, TCP and calcium carbonate; and TTCP and DCPA.

However, in an embodiment of the invention, the cement powder does not comprise substantial amounts of carbonate. Preferably the cement powder comprises less than 10 wt.% of carbonate containing compounds, more preferably less than 5 wt.%, more preferably less than 3 wt.%, more preferably less than 2 wt.%, more preferably less than 1 wt.%, more preferably less than 0.5 wt.%, most preferably less than 0 1 wt.%. In an embodiment, the cement powder is substantially or entirely free from carbonate compounds.

In a preferred embodiment, the cement powder according to the invention comprises at least 50 wt.%, more preferably at least 70 wt.%, still more preferably at least 80 wt.%, still more preferably at least 85 wt.% of the calcium phosphate component, e.g. approximately 90 wt.%.

In accordance with the present invention, the bone cement powder comprises a pore forming agent based on Glucono delta-lactone (GDL).

GDL is a cyclic ester of gluconic acid. Glucono delta-lactone (GDL) is a GRAS (Generally Recognized as Safe) food additive as recognized by the FDA. GDL is generally used as curing, pickling agent or leavening agent, pH controller and sequestrant. Its end- product gluconic acid can be naturally eliminated by the body since it is a metabolite that occurs after oxidation of glucose.

The term 'pore forming agent' is used to indicate that the agent upon introduction of the cement in the physiological environment gives rise to the formation of interconnected macropores in the CPC material, thus conferring macroporosity. A "macropore" is a pore with a diameter of above 100 μηι. Thus "macroporosity" describes a state of the cement where interconnected macropores are present. The macroporosity of the implant allows infiltration by cells, bone formation, bone remodeling, osteoinduction, osteoconduction, and/or faster degradation of the set CPC. Pores in the set CPC are thought to improve the osteoinductivity or osteoconductivity of the composite by providing space for cells such as osteoblasts, osteoclasts, fibroblasts, cells of the osteoblast lineage, stem cells, etc. The pores facilitate ingrowth of bone tissue. Pores in the composite may also provide for easier degradation of the composite as bone is formed and/or remodeled.

As explained herein before, GDL can confer macroporosity by virtue of the fact that it produces acid at high rate upon introduction into the physiological environment, so that the size and shape in which it is presented in not critical. Hence, the invention entails embodiments wherein the GDL is included in the bone cement or bone cement powder in the form of a fine powder, a particulate material, a granulate, etc. In an embodiment of the invention GDL is included in the bone cement or bone cement powder in the form of microparticles, microspheres or microbeads, which may be spheroidal, cuboidal, rectangular, elonganted, tubular, fibrous, disc-shaped, platelet-shaped, polygonal, etc. In certain embodiments, the GDL based pore forming agent is a material containing particles with a diameter ranging from approximately 20 μιη to approximately 800 μηι. Preferably, the diameter of the GDL microparticles or microbeads is between 50 and 300 μπι, preferably 80 and 250 μιη, more preferably 100 and 200 μηι. GDL is commercially available as dense particles of approximately 300 μιτι in size. It is believed that after embedding these GDL particles in CPC, the particle size is somewhat reduced. As will be shown in the examples, the use of this material resulted in macroporosity optimal for new bone tissue ingrowth. These material are available commercially, e.g. from Merck. Furthermore, it will be apparent to those skilled in the art how to prepare GDL material in the form of suitable microspheres or microbeads starting from commercially available materialsusing conventional techniques.

In one embodiment of the invention the pore forming agent mainly comprises of GDL, meaning that it comprises more than 95 wt.%, more than 97.5 wt.%, more than 99 wt.%, more than 99.5 wt.% or more than 99.9 wt.% of GDL. In one embodiment the pore forming agent essentially consists of GDL. Embodiments are however also envisage wherein the GDL based pore forming agent comprises particles comprising a GDL core in an outer coating which is degradable in the physiologic environment, e.g. a PLGA coating. Yet, in another embodiment of the invention the GDL based pore forming agent may comprise additional pore forming agents, such as PLGA microparticles or microbeads, suitable examples of which are know from the prior art. It will be appreciated by those skilled in the art, that such embodiments allow for further control and adjustment of the process of macroporosity formation after injection and setting of the CPC of the invention. Typically, in these embodiment of the invention the GDL based pore forming agent comprises at least 30 wt.%, at least 50 wt.%, at least 60 wt.%, at least 70 wt.%, or at least 80 wt.% of GDL. The amount of the GDL based pore forming agent may vary in the bone cement powder between about 0.1 and about 50%, preferably between about 0.5 and about 30%, more preferably between about 1 and about 20%, most preferably between about 2 and about 10% by weight of the total amount of the cement powder according to the invention.

In an embodiment, the ratio between the calcium phosphate compound and the GDL pore forming agent is from 3 : 1 to 50: 1 and more preferably from 4: 1 to 20: 1, preferably 5 : 1 to 10: 1, depending on the desired porosity.

In an embodiment, the ratio (w/w) between the GDL pore forming agent and carbonate is 1/1 or more, more preferably 2/1 or more, more preferably 5/1 or more, more preferably 10/1 or more, more preferably 50/1 or more, most preferably 100/1 or more. In an embodiment, the bone cement powder is substantially or entirely free from carbonate.

A second aspect of the invention is a bone cement resulting from the mixing between a cement powder as described in the foregoing and a liquid phase. The term "bone cement" means a flowable and/or moldable viscous material that can harden into a solid mass that can be used to fill bone voids or fissures and attach to and/or structurally engage local bone structure. The cement typically comprises a liquid component and the bone cement solids dispersed within said liquid component. The composition comprising the powder and the liquid may typically be in the form of a paste or semi-solid.

An appropriate liquid phase includes one or more of the following: deionized water, dilute phosphoric acid, dilute organic acids (acetic, citric, succinic acid), sodium phosphate, sodium carbonate or bicarbonate, sodium alginate, sodium bicarbonate, sodium chondroitin sulphate. Deionized water and aqueous sodium phosphate solution, especially Na2HPC>4 solution and Na ₂ HP0 ₄ /NaH ₂ P0 ₄ aqueous solution, are preferred. . For example, a solution of 2 to 3% by weight of a ₂ HP0 ₄ in distilled water can be used.

The pH of the liquid phase is preferably between 5 to 10, more preferably between 5 and 9, most preferably between 5 and 7.

Preferably, the liquid phase/solid phase (L/S) ratio is between about 0.25 and about 0.7 ml/g, more preferably between about 0.3 and about 0.6 ml/g, the most preferably is about 0.4 ml/g or about 0.5 ml/g.

The setting time, which can range from about 10 to about 60 min, preferably about 10 to about 30 min, depends on the composition of the powder and liquid components, the powder-to-liquid ratio, proportion of the calcium phosphate components and the particle sizes of the powder components. The setting time of the cement is an important property of the cement. If the setting time is too fast, the surgeon does not have time to use the cement before it is hard. If the setting time is too long, the surgeon must wait until he/she can close the wound. For the purposes of the present invention, setting time will be defined as the time required for the composition to reach a set or solid state after being applied in a fluid or paste state. In particular, setting time refers to the time required for the composition to attain a specified degree of rigidity. A skilled person would realize that for some applications, a shorter setting time is desired, whereas for other applications, a longer setting time is desired. For example, in orthopedic surgical applications, a long setting time may be required to allow the surgeon to have sufficient time to load and/or inject the composition to the appropriate bone site of the patient. The setting time of the composition according to any aspect of the present invention, under physiological conditions, may be from 1 minute to 1 hour; from 4 minutes to 45 minutes, from 5 minutes to 40 minutes, from 6 minutes to 35 minutes, from 7 minutes to 30 minutes, from 8 minutes to 25 minutes, from 9 minutes to 20 minutes, from 10 minutes to 15 minutes, from 11 minutes to 13 minutes. Even more in particular, the setting time is from 4 minutes to 7 minutes.

Another important characteristic is inj ectability, meaning that preferably the cement is flowable or syringable at ambient temperature. More preferably, the cement is sufficiently fluid to flow through a needle with a diameter of a few millimeters, preferably between 1 and 5 mm, typically at ambient temperature. In a preferred embodiment of the invention the cement has a viscosity of from 25 to 1000 mPas and in some embodiments, preferably 50 to 1000 mPas, measured at 20 °C.

Preferably, the bone cement of the invention is "Biocompatible", meaning that it does not elicit detrimental effects associated with the body's various protective systems, such as cell and humoralassociated immune responses, e. g. , inflammatory responses and foreign body fibrotic responses. The term biocompatible also implies that no specific undesirable cytotoxic or systemic effects are caused by the material when it is implanted into the patient.

Furthermore, in most embodiments, the cement is sterile. Sterilization of the cement can be accomplished using conventional techniques such as gamma-irradiation, sterility filters or by using strictly sterile procedures.

The bone cement or bone cement powder as described in the previous sections may suitably comprise at least one bioactive agent, i.e. in addition to the calcium phosphate cement and the glucono-delta-lactone material. For the purposes of the present invention the term, "bioactive agent" generally refers to a molecule that cause(s) a biological effect when administered in vivo to mammals, especially humans and which can be of any nature. For instance synthetic organic compounds, peptides, proteins, carbohydrates (including monosaccharides, oligosaccharides, and polysaccharides), steroids, nucleic acids, nucleotides, nucleosides, oligonucleotides, genes, lipids and hormones may be incorporated in the present bone cement.

In one particularly preferred embodiment of the invention said bioactive agent is an osteogenic agent. An "osteogenic agent," as used herein, generally refers to an agent capable of inducing and/or supporting the formation, development and growth of new bone, and/or the remodeling of existing bone. Particularly preferred examples include osteogenic agents. The osteogenesis process typically involves the deposition of new bone by cells called osteoblasts. Many osteogenic agents function, at least in part, by stimulating or otherwise regulating the activity of osteoblast and/or osteoclasts.

Suitable osteogenic materials may include a growth factor that is effective in inducing formation of bone. Desirably, the growth factor will be from a class of proteins known generally as bone morphogenic proteins (BMPs), and may in certain embodiments be recombinant human (rh) BMPs. and any one of the many known bone morphogenic proteins (BMPs), including but not limited to BMP-1, BMP-2, BMP-2A, BMP-2B, BMP-3, BMP-3b, BMP-4, BMP-5, BMP-6, BMP-7, BMP-8, BMP-8b, BMP-9, BMP-10, BMP-11, BMP-12, BMP-13, BMP-14, BMP-15 as well as recombinant forms of any one of the above BMP's. In a preferred embodiment a human BMP is used, preferably a recombinant human BMP.

Other therapeutic growth factors or substances may also be used in the bone cement or bone cement powder of the invention, to stimulate bone formation, such as platelet-derived growth factors, insulin-like growth factors, cartilage-derived morphogenic proteins, growth differentiation factors, such as GDF-5, and transforming growth factors, including TGF-a and TGF-β, extracellular matrix-associated bone proteins such as, osteocalcin, bone sialoprotein (BSP) and osteocalcin in secreted phosphoprotein (SPP)-l, type I collagen, fibronectin, osteonectin, thrombospondin, matrix-gla-protein, SPARC, osteopontin, osteogenin, osteoinductive factor (OIF), basic Fibroblast Growth Factor (bFGF), acidic Fibroblast Growth Factor (aFGF), vascular endothelial growth factor (VEGF), Growth Hormone (GH), and osteogenic protein- 1 (OP-1).

The amount of osteogenic substance useful herein is that amount effective to stimulate increased osteogenic activity of infiltrating progenitor cells, and will depend upon several factors including the size and nature of the defect being treated, and the device and particular protein being employed.

The bone cement or bone cement powder may also comprise progenitor and/or stem cells derived from embryonic or adult tissue sources and/or taken from culture. Illustratively, compositions of the invention may incorporate cells derived from blood, bone marrow, or other tissue sources from the patient to be treated (autologous cells) or from a suitable allogenic or xenogenic donor source. Living microorganisms, plant cells, animal cells or human cells may be incorporated in the cement as well. The cells may be selected from the group consisting of osteoblasts, osteocytes, osteoclasts, osteoprogenitor cells, (bone tissue), chondrocytes (articular cartilage), fibrochondrocytes (meniscus), ligament fibroblasts (ligament), skin fibroblasts (skin), tenocytes (tendons), myofibroblasts (muscle), mesenchymal stem cells and keratinocytes (skin).

Quite often, bone defects are not due to a traumatic event, but to a disease, e.g. bone tumour, infection, etc. In some cases, it may be advantageous to incorporate other drugs in the cement, such as antimicrobials, antibiotics, antimyobacterial, antifungals, antivirals, antineoplastic agents, antitumor agents, agents affecting the immune response, blood calcium regulators, general inhibitors of the allergic response, antihistamines, local anesthetics, nonsteroidal anti-inflammatory agents, and steroidal anti-inflammatory agents. Owing to their structure and their dissolution property, the calcium phosphate cements are able to slowly release the active ingredients into the environment within a few days after implantation.

I view of the risk of extravasation of the cement into the tissues surrounding bone, it may be desirable to visualise the cement. The easiest way is to increase the radio-opacity of the cement, for example by means of contrasting agents. Any suitable radiopaque material may be used. The radiopaque material may be an oxide or halogen salt of a heavy metal. Examples of radiopaque materials include, but are not limited to, barium sulfate, tungsten, bismuth compounds, tantalum, zirconium, platinum, gold, silver, stainless steel, titanium, alloys thereof, combinations thereof, or other equivalent materials for use as radiographic agents.

In accordance with this invention, other additives may be included in the bone cement or bone cement powder as defined in any of the foregoing, e.g. to adjust setting times, increase injectability, reduce cohesion or swelling time, etc. For example, the composite may further include one or more of an initiator, accelerator, catalyst, solvent, wetting agent, lubricating agent, labeling agent, plasticizer, etc.

A very efficient way to accelerate the setting time is to have large concentrations of phosphate ions in the mixing solution, e.g. by adding a soluble phosphate salt to the bone cement powder. A soluble phosphate salt may, alternatively be pre-dissolved in the liquid phase. Another way to accelerate the setting reaction is to add germs for apatite crystal growth, as the nucleation step of the setting reaction is a limiting factor. Typically, apatite crystals can be used, preferably a calcium-deficient hydroxyapatite or hydroxyapatite powder. Small amounts (a few weight percents) are sufficient to drastically reduce the setting time.

When the setting time is too short, various setting additives can be added to increase the setting time. Typical examples are compounds which inhibit the nucleation and/or growth of apatite crystals. Common examples are pyrophosphate, citrate or magnesium ions. One particularly interesting compound is calcium carbonate. Calcium pyrophosphate (CPP), calcium sulfate dihydrate (CSD) and calcium sulfate hemihydrate (CSH) can also increase the setting time.

The bone cement or bone cement powder of this invention may further comprise a binder. The binder provides enhanced viscosity and cohesiveness of the composition, allowing the surgeon to position and shape the composition within the voids, defects or other areas in which new bone growth is desired. The enhanced cohesiveness of the composition also prevents particle migration associated with grafting materials for orthopedic, maxillofacial and dental applications. The minimum amount of binder is that amount required to give easy formability and provide sufficient particle cohesion and shape retention during the period of tissue ingrowth. The binders contemplated as useful herein include, but are not limited to suspending agents, viscosity-producing agents, gel-forming agents and emulsifying agents. More preferably, the binder is selected from the group consisting of sodium alginate, hyaluronic acid, methylcellulose, carboxy methylcellulose, carboxy methylcellulose sodium, carboxy methylcellulose calcium, hydroxypropyl methylcellulose, hydroxybutyl methylcellulose, hydroxyethyl methylcellulose, hydroxyethylcellulose, methylhydroxyethyl cellulose, hydroxyethyl cellulose and admixtures thereof.

Another aspect of the invention provides kits for producing a bone cement of the invention. One embodiment provides a kit of parts comprising a first and a second reservoir, each holding distinct compositions containing one or more substances selected from calcium phosphate compound, GDL based pore forming agent and a liquid phase, all as described in any of the foregoing. Preferably, said first reservoir holds the bone cement powder, and said second reservoir holds the liquid phase. Either one or both of said first and second reservoir may contain additional agents such as the bioactive agent, radiopaque material, binder, setting additives, etc. In some embodiments, the kits comprise an additional reservoir holding one or more of these additional agents.

The compositions in the first and second reservoir, upon admixing, should yield the bone cement described above. This means, as will be understood by those skilled in the art, that the amounts, concentrations and relative ratios of the components in each composition and in each reservoir should be chosen so as to obtain the values defined above in the final bone cement, i.e. after admixing the individual components and/or compositions provided in the kit.

In some embodiments, the kits comprise a mixing device. In some embodiments, the kits comprise a dispensing device. The dispensing device may be an applicator in communication with a mixing device and/or a reservoir adapted for containing the cement. In some embodiments, the dispensing device comprises a catheter. In some embodiments, the dispensing device comprises a syringe. In accordance with the invention the afore described reservoirs may be separate reservoirs or they may be compartments of an integrated reservoir system or dispensing device. In some embodiments a single device is provided comprising a first and second reservoir as described above and a means for mixing the compositions from the reservoirs, as well as an applicator. Kits preferably contain instructions for use.

The bone cements described herein and, hence, the kits for preparing them, are useful in reconstructive treatment or intervention, such as orthopedic, craniofacial, oral and maxillofacial treatment. The cement mixture can be used as a joiner or filler for the assembly of bone and/or cartilage tissue surfaces, which are not in direct contact, and for binding bone or cartilage tissue to metallic or synthetic prosthetic devices.

The bone cement may be used in particular as a bone void filler. Bone fractures and defects, which result from trauma, injury, infection, malignancy or developmental malformation can be difficult to heal in certain circumstances. If a defect or gap is larger than a certain critical size, natural bone is unable to bridge or fill the defect or gap. There are several deficiencies that may be associated with the presence of a void in a bone. The bone void may compromise the mechanical integrity of the bone, making the bone potentially susceptible to fracture until the void becomes ingrown with native bone. Accordingly, it is of interest to fill such voids with the CPC of the present invention, which allows the void to eventually fill with naturally grown bone. Open fractures and defects in practically any bone may be filled with composites according to various embodiments without the need for periosteal flap or other material for retaining the composite in the fracture or defect.

Many orthopedic, periodontal, neurosurgical, oral and maxillofacial surgical procedures require drilling or cutting into bone in order to harvest autologous implants used in the procedures or to create openings for the insertion of implants. In either case voids are created in bones. Surgically created bone voids may be filled using a bone cement of the invention The CPC according the invention can be used for dental applications. Potential dental applications are: repair of periodontal defects, sinus augmentation, maxillofacial reconstruction, pulp-capping materials, cleft-palate repair, and as adjuvants to dental implants.

In an aspect of the invention, a method is provided for reconstructing or repairing a defect or cavity in a human or mammalian bone or cartilage tissue comprising filling the defect or cavity with a bone cement as previously described. The bone cement may be introduced at the site of the bone void or defect using any technique known in the art. The present methods may include a surgery step but bone cements according to the invention can get to inaccessible parts of the body with minimally invasive surgery procedures to reduce damage and pain while hastening return to function. Hence, in a preferred embodiment of the invention a method of aesthetic or reconstructive intervention in a subject is provided, said method comprising i) providing a bone cement as defined in any one of claims 7-9; ii) applying the bone cement to a defect or cavity in a tissue of said subject so as to fill said defect or cavity; and iii) allowing the bone cement to set. This method of treatment comprises the introduction of an injectable CPC according to the invention into the bony defect or fracture through a needle The bone cement may be introduced in the body by injection or endoscopic administration using techniques and equipment generally known by those skilled in the art. Injection of the cement may be limited by the viscosity thereof which controls the injectability or syringeability of the solutions. A needle having a gauge of 20 or below is ideal for injection of bone cement.

According to one embodiment of the invention the bone cement is prepared only shortly before using it in a method of the invention. Typically, this will be accomplished by mixing a bone cement powder and a liquid phase as described herein before.

Composites of the present invention can be used as bone void fillers alone or in combination with one or more other conventional devices, for example, to fill the space between a device and bone. Examples of such devices include, but are not limited to, bone fixation plates (e.g., cranofacial, maxillofacial, orthopedic, skeletal, and the like); screws, tacks, clips, staples, nails, pins or rods, anchors (e.g., for suture, bone, and the like), scaffolds, scents, meshes (e.g., rigid, expandable, woven, knitted, weaved, etc), sponges, implants for cell encapsulation or tissue engineering, drug delivery (e.g., carriers, bone ingrowth induction catalysts such as bone morphogenic proteins, growth factors, peptides, antivirals, antibiotics, etc), monofilament or multifilament structures, sheets, coatings, membranes (e.g., porous, microporous, resorbable, etc), foams (e.g., open cell or close cell), screw augmentation, cranial, reconstruction, and/or combinations thereof. In any of the above methods, the subject is typically a mammal. In certain preferred embodiments, the subject is a human. In other embodiments, the subject is a domesticated animal such as a dog, cat, horse, etc.

Another aspect of the present invention concerns the use of glucono-delta-lactone as a pore forming agent in a calcium phosphate cement.

It is contemplated that any embodiment discussed in this specification can be implemented with respect to any of the methods, objects, compositions and uses of the invention, and vice versa. The principal features of this invention can be employed in various embodiments without departing from the scope of the invention. Those skilled in the art will recognize, or be able to ascertain using no more than routine experimentation, numerous equivalents to the specific procedures described herein. Such equivalents are considered to be within the scope of this invention and are covered by the claims.

Furthermore, for a proper understanding of this invention and its various embodiments it should be understood in this document and the appending claims, the verb "to comprise" and its conjugations are used in its non-limiting sense to mean that items following the word are included, but items not specifically mentioned are not excluded. In addition, reference to an element by the indefinite article "a" or "an" does not exclude the possibility that more than one of the element is present, unless the context clearly requires that there be one and only one of the elements. The indefinite article "a" or "an" thus usually means "at least one".

The following examples describe various new and useful embodiments of the present invention. These examples are for illustrative purposes and do not limit the scope of the invention. Examples

Experiment I: Accelerated calcium phosphate cement degradation due to incorporation of glucono delta-lactone as porogen Background

The aim of the current investigation was to study the effects of GDL as CPC porogen in vitro and in vivo, to evaluate this novel material combination and to compare its in vivo performance as bone filler materials to another hydrolytically degradable composite (CPC- PLGA) and an enzymatically degradable composite (CPC-gelatin). The main hypothesis was that GDL would degrade in a fast rate creating CPC porosity at a very early implantation stage and therefore allowing bone substitution of the material at initial time points.

Two different CPC-GDL percentages were studied, one with a minimal GDL amount (10 %) to investigate bone reaction to this novel porogen material and another one (30 %) with a similar macroporosity to the other tested composites of different nature. To this end, four types of CPC composites (CPC-gelatin, CPC-PLGA, CPC-10%GDL, CPC-30%GDL) embedded in CPC were implanted in cylindrical defects as created in the femoral condyle of New Zealand White rabbits. CPC degradation and new bone formation were evaluated histologically and histomorphometrically after 2 and 12 weeks of implantation.

Materials & Methods

CPC powder consisted of a mixture of 85% alpha-tricalcium phosphate ( -TCP; CAM Bioceramics BV, Leiden, the Netherlands), 10% dicalcium phosphate anhydrous (DCPA; Sigma Aldrich, St. Louis, MO, USA) and 5% precipitated hydroxyapatite (pHA; Merck, Darmstadt, Germany). Na ₂ HPC>4 was purchased from Merck (Darmstadt, Germany).

Poly(lactic-co-glycolic acid) (PLGA) Purasorb ^® materials were obtained from Purac Biomaterials (Gorinchem, the Netherlands). Purasorb ^® PDLG 5002A (Mw = 17 kDa, acid terminated, L:G = 50:50) was used to prepare microspheres. Poly vinyl alcohol (PVA; 88% hydrolyzed, MW 22000) was obtained from Acros (Geel, Belgium) and isopropanol (IPN; analytical grade), dichloromethane (DCM; analytical grade) and D-(+)-Glucono-delta-latone were obtained from Merck (Darmstadt, Germany).

Anionic gelatin (from bovine skin, Type B Bloom 225, IEP-5) as well as acetone and glycine was purchased from Sigma-Aldrich (St. Louis, MO, USA). Olive oil and glutaraldehyde (GA, 25 wt% solution in water) were commercially available from Acros Organics (Geel, Belgium).

Microspheres preparation and characterization

Dense PLGA microspheres were prepared by a single emulsion technique. Briefly, 0.2 g of PLGA was dissolved in 2 mL of DCM in a 20 mL glass tube. Then, this solution was transferred into a stirred beaker containing 100 mL of 0.3% PVA solution. Subsequently, 50 mL of 2% IPN solution was added. The solution was stirred for one hour. The microspheres were allowed to settle for one hour and the clear solution was decanted. The remaining suspension was centrifuged and the clear solution on top was aspirated. Finally, the microspheres were frozen, freeze-dried for 24 hours and stored at -20°C.

The morphology and size distribution of the PLGA microspheres was determined by light microscopy. Microspheres were suspended in H ₂ O and pictures were taken with an optical microscope equipped with a digital camera (Leica/Leitz DM RBE Microscope system, Leica Microsystems AG, Wetzlar, Germany). The size distribution of the microspheres was determined by digital image software (Leica Qwin ^® , Leica Microsystems AG, Wetzlar, Germany) using a sample size of at least 300 microspheres. In addition, morphology of the microspheres was observed by Scanning Electron Microscopy (SEM, JEOL 6310) at an accelerating voltage of 10 kV.

Gelatin microspheres were prepared using a surfactant-free water-in-oil emulsion method. A 10 w/v% gelatin solution was obtained by dissolving gelatin in deionized water at 45°C. Thereafter, 10 mL of this gelatin solution was added dropwise into a 500ml round bottom flask containing 300 mL olive oil while stirring at 500 rpm (upper stirrer) at 45°C. Spontaneous gelation of the gelatin droplets was then obtained by moving the emulsion into an ice bath under continuous stirring. After 30 min, 150ml of chilled acetone (4°C) was then added to the emulsion followed by another 15 min of stirring. Gelatin microspheres without crosslinking were collected by filtration (Whatman 90 mm filterpaper) and washed with chilled acetone to remove residual olive oil. Subsequently, gelatin microspheres were re- suspended in water/acetone (1/3 in volume ratio) solution, and further crosslinked using GA with a molar ratio of GA relative to the amount of free amine groups present in gelatin ([ H ₂ ] gelatin ) of 0.5. After crosslinking for 16 h, glycine solution (100 mM ) was added into the microsphere suspension to block unreacted aldehyde groups from GA. The suspension was then subjected to three cycles of centrifugation (5000 rpm for 5 min) and re- suspension in deionized water by vortexing. After freeze-drying, microspheres were stored at 4 °C until further use.

GDL in vitro studies

Different CPC-GDL porosities were generated by the use of variable volumes of the CPC liquid phase (2% Na ₂ HP0 ₄ ). For CPC-30%GDL 230 was the selected volume in order to obtain a composite of approximately 50% porosity (to be comparable with the CPC- PLGA and CPC-gelatin composites) (Figure 1). For CPC-10%GDL 290 μL· was the volume of choice since it results in an paste with similar injectability and setting properties as when CPC-30%GDL with 230 μΕ is used. Different volumes of 2% Na2HPC>4 solution were added to the CPC composites (350 μL to PLGA and gelatin composites and 290 μΐ,, 210 μL to CPC-10%GDL and CPC- 30%GDL respectively). Subsequently, the syringes were mixed vigorously for 30 seconds and the cement was injected into Teflon molds (3 mm x 6 mm). Pre-set composites containing either PLGA, GDL or gelatin were scanned using a uCT (Sky scan -1072 X-ray microtomograph, TomoNT version 3N.5, Skyscan ^® , Bel gium; X-ray source was set to 100 kV, current to 98 μΑ and the resolution was 7 μιη pixel). Cone beam reconstruction was performed and the data were analyzed by CT analyzer (version 1.4, Skyscan ^® ).

To investigate the hydrolitical rate of GDL, pre-set composites were incubated at 37 °C in 1.5 mL of PBS for 30 minutes, 1 hour, 2 hours, 1 day, 3 days and 5 days. The amount of gluconic acid in the incubation media of the different CPC-GDL composites was analyzed by Reverse Phase High Performance Liquid Chromatography (RP-HPLC) as a measurement of GDL degradation. The system consisted of a Hitachi L2130 HPLC pump, Hitachi L-2400UV detector and Hitachi L-2200 auto sampler, and an Atlantis dC18 column (Waters) 250 mm x 4.6 mm, 5 μηι. Two mobile phases were used; 1% acetonitrile in 20 mM aH ₂ P04 pH= 2.2 (mobile phase A) and 100% acetonitrile (mobile phase B). The flow was 0.5 mL per minute, the injection volume 40 μL and the UV detection wavelength 210 nm.

The chemical composition of the pre-set CPC-GDL samples was evaluated crushing the samples with a mortar until the discs were reduced to powder and then analyzing them using x-ray diffraction (XRD) (PW3710 Philips, The Netherlands) was performed with a Cu K _a radiation source with a wavelength of 1.5405 A at a voltage of 40 kV and a current of 30 mA. Spectra were collected for 2Θ values of 3°-45° in continuous mode. The positions and intensities of the XRD peaks were used to identify the underlying structure of the CPC materials.

Preparation and characterization of injectable CPC composites for in vivo implantation

In order to prepare CPC-PLGA composites, 0.6 g of dense PLGA microspheres was added to 0.8 g of CPC powder inside a 2 mL plastic syringe.

Two different CPC-GDL composites were prepared containing either 10% (w/w) GDL or 30% (w/w) GDL, for this propose 0.1 g of GDL were mixed with 0.9 g of CPC (CPC- 10%GDL) or 0.3 g of GDL were mixed with 0.7 g of CPC (CPC-30%GDL) both in a 2 mL plastic syringe.

Finally, three different CPC compositions were created containing three types of porogen materials: (1) PLGA microspheres, (2) 10% GDL (3) 30% GDL. All syringes were sealed with a closed tip and sterilized using gamma radiation (25-50 kGy; Isotron BV, Ede, the Netherlands).

Gelatin microspheres were sterilized by ethylene oxide in order to avoid possible cross linking effects of the gamma radiation over the gelatin microspheres (Isotron BV, Ede, the Netherlands). Prior to implantation the CPC-gelatin composite was prepared under sterile conditions by adding 200 of dH20 to 0.05 g of gelatin microspheres, let them swell for 10 minutes and mix them with 0.95 g of CPC powder (previously sterilized by gamma radiation) in a 2 mL plastic syringe.

Before implantation, different volumes of sterile 2% Na2HP04 solution was added to each sample and mixed vigorously for 30 sec (Silamat ^® mixing apparatus, Vivadent, Schaan, Liechtenstein) in order to achieve the most optimal injectable properties for each composite. Therefore 350 μΐ _^ of NaiHPC solution were added to gelatin and PLGA composites, whereas 290 μΐ. and 210 μΐ. were added to CPC-10%GDL and CPC-30%GDL respectively. Surgical procedure

24 healthy skeletally mature female 8 weeks old New Zealand White rabbits with a weight between 3.2 and 3.6 kg were used as experimental animals. The protocol was approved by the Animal Ethical Committee of the Radboud University Nijmegen Medical Centre (Approval no: RU-DEC 2010-174) and national guidelines for the care and use of laboratory animals were applied.

Surgery was performed under general inhalation anesthesia. The anesthesia was induced by an intravenous injection of Hypnorm (0.315 mg/mL fentanyl citrate and 10 mg/ml fluanisone) and atropine, and maintained by a mixture of nitrous oxide, isoflurane and oxygen through a constant volume ventilator. To reduce the peri-operative infection risk, the rabbits received antibiotic prophylaxis (Baytril®, 2.5% (Enrofloxacin), 10 mg/kg). The animals were immobilized on their back and the hind limbs were shaved, washed and disinfected with povidone-iodine. After exposure of the distal femoral condyle a 1.0 mm pilot hole was drilled. The hole was gradually widened with drills of increasing size until a final defect size of 4 mm in width and 6 mm in depth was reached. Low rotational drill speeds (max. 450 rpm) and constant physiologic saline irrigation were used. After preparation, the defects were thoroughly irrigated and packed with sterile cotton gaze to stop bleeding. Surgery was performed in both legs of the rabbits and one defect was created in each condyle.

The four different CPC formulations were injected into the created condylar defects in a randomized manner (N=6). After injection, the superfluous cement was removed with a scalpel blade and soft tissues were closed layer-by-layer using resorbable Vicryl sutures after approximately 10 minutes of setting.

After 2 and 12 weeks of implantation, rabbits were euthanized using an overdose of Nembutal® and the femoral condyles were harvested for evaluation.

Histological procedures

After harvesting of the femoral condyles and removal of surrounding soft tissues, each condyle was divided into medial and lateral halves along its longitudinal axis. Subsequently, the medial half was fixed in 4% formaldehyde for 2 days, dehydrated in a graded series of ethanol and embedded in methylmethacrylate (MM A). After polymerization, at least three 10 μιτι sagittal cross sections of the condyle part containing the composite implants were prepared using a sawing microtome technique. Sections were stained with methylene blue/basic fuchsin and examined with a light microscope (Leica Microsystems AG, Wetzlar, Germany).

Histological and -morphometrical analysis

In addition to a subjective histological description, quantitative histomorphometrical analysis was performed. Therefore, a 4 mm diameter circular area was superimposed on the bone defect area (Figure 2). Digital images (magnitude 5x) were recorded and bone as well as cement content were determined by color recognition using image analysis techniques (Leica Qwin Pro-image analysis system, Wetzlar, Germany).

Statistical analysis

All statistical analyses were performed with SPSS ^® software (IBM Corporation, Somers, NY, USA). Significant differences between the groups were determined using analysis of variance (ANOVA). A Tukey-Kramer (Multiple Comparisons) Test was applied for comparison of the groups at different implantation periods. Results were considered significant if p<0.05. Results

Physicochemical characterization of porogens and composites

Microspheres sizes were 42 ± 5 μπι for PLGA microspheres and 37 ± 10 μιη for gelatin microspheres. In contrast, GDL particles size was equal to 595 ± 40 μπι (Figure 3). SEM micrographs of the pre-set CPC composites revealed a more compact and less nanoporous structure in CPC-30%GDL than in the rest of the studied composite materials (Figure 4).

Initial CPC porosity due to the porogens present in pre-set composites was 53 ± 3 % for CPC-gelatin, 50 ± 4 % for CPC-PLGA, 19 ± 2 % for CPC-10%GDL and 48 ± 2 % for CPC-30%GDL.

CPC-GDL in vitro tests

The rapid GDL dissolution rate when embedded in CPC can be observed in the RP- HPLC results (Figure 5). After 1 day of incubation in PBS the majority of the GDL was released and detected by the HPLC (96 % ± 2).

Regarding the XRD results, the CPC phases were similar with or without the addition of GDL indicating that GDL does not change the original CPC structure. Extra XRD peaks can be observed in the samples containing GDL, after analysis of the GDL powder it can be concluded that these peaks are representative of the GDL structure, as discernable by comparing both XRD patterns (Figure 6).

Clinical observations

All the 24 rabbits exhibited good health and did not show any wound complication. At the end of the two implantation periods, a total of 48 implants were retrieved (12 rabbits x 24 implants after 2 weeks' implantation and 12 rabbits x 24 implants after 12 weeks' implantation). At retrieval, no visual signs of inflammatory or adverse tissue reactions were observed.

Descriptive light microscopy

A uniform tissue reaction was seen between the specimens of the same group and implantation time. Representative light microscope sections corresponding to 2 and 12 weeks implantation time are presented in Figure 7 and Figure 8, respectively.

Examination of the sections after 2 weeks implantation revealed variable amounts of cement degradation and newly formed bone in the ROI. The newly formed bone displayed cancellous structure similar to the preexistent bone in the vicinity of the defect location. CPC- PLGA composites presented intact round morphology indicating no degradation of the CPC material and therefore no bone tissue formation in the ROI. Similar results can be observed for the CPC-gelatin composites where no degradation of the CPC can be detected. In contrast, CPC degradation can be seen in the defects containing CPC-10%GDL, initial degradation of the material occurred in the peripheral area of the round defect and newly formed bone can be observed. For CPC-30%GDL containing materials some minor degradation could be observed in the peripheral areas of the defect, however big non dissolved GDL crystals could be found within the CPC matrix

After 12 weeks of implantation, the majority of the CPC-PLGA composites were completely degraded and only small remains of CPC were found in between the newly formed interconnected bone network. CPC-gelatin materials undergo some degradation of the implant peripheral area and new bone formation occurred throughout the degraded areas. Regarding CPC-10%GDL composites, major degradation of the material was detected and new bone formation can be observed in the ROI. Similar to the CPC-PLGA, after CPC- 10%GDL degradation a cancellous bone network can be observed in the defect location. In CPC-30%GDL composites degradation of the outer regions of the circular implant were found, newly formed bone was observed in the areas where the material was degraded.

At both time points, when new bone formation occurs in close contact with the implanted material with no signs of inflammatory reaction were be observed.

Bone formation

Quantitative evaluation of new bone formation within the ROI revealed that the amount of newly formed bone varied considerably among the experimental groups (Figure 9). After 2 weeks implantation CPC-10%GDL showed significantly higher new bone formation (32.8 %) compared to all other groups (p<0.001). CPC-30%GDL presented significantly more newly formed bone (12.5 %) than composites with PLGA (2.0 %) or gelatin (1.9 %) microspheres (p<0.05).

After 12 weeks implantation CPC-PLGA (64.0 %) and CPC-10%GDL (66.0 %) revealed significantly (p<0.001) enhanced bone formation in comparison to CPC-30%GDL (25.8 %) and CPC-gelatin (4.4 %). CPC-30%GDL composites displayed significant more bone formation than CPC-gelatin (p<0.01).

Bone formation in relation to 100% trabecular bone

Trabecular bone formation in relation to a ROI filled by 100 % bone was calculated. After 2 weeks implantation results revealed 46.8 %, 17.9%, 2.9 %, 2.8 % new bone formation CPC-for samples in which CPC-10%GDL, CPC-30%GDL, CPC-PLGA and CPC-gelatin were the injected composites respectively.

After 12 weeks implantation both CPC-PLGA (91.4 %) and CPC-10%GDL (94.3 %) specimens revealed no significant differences in comparison to a ROI filled with trabecular bone. However significant less bone formation occurred when CPC-30%GDL (36.8 %) and CPC-gelatin composites were used (6.2 %) (p< 0.01). (Figure 10).

Discussion

The aim of this study was to evaluate GDL as a CPC porogen both in vitro and in vivo as injectable bone filler in femoral condyle defects in New Zealand White Rabbits. The hypothesis was that GDL could induce fast generation of CPC porosity and therefore enhance CPC degradation and subsequent bone ingrowth within the implant material at an early stage. CPC-GDL composites performance in the in vivo situation was compared to two other porogens materials; PLGA and gelatin microspheres. The results of the present study demonstrated that GDL led to enhanced CPC degradation at early time points (2 weeks) and practically complete resorption of the material and substitution by newly formed bone tissue at later time points (12 weeks).

The initial GDL concentration affects the initial pore size, leading to smaller pores when a smaller initial GDL content is used. Since the ideal porosity and pore size of CPC materials remains unclear; GDL presents the advantage of a large initial size (595 ± 40 μπι) which can be tailored by the addition of different volumes of 2% Na2HPC>4 leading to composites with different initial macroporosity percentages and pore size.

The PLGA microspheres used in this study were fabricated from 17 kDa acid terminated polymer and with dense morphology. Microspheres of these chemical and morphological characteristics were the material of choice due to its previously proven fast degradation and new bone formation in combination with injectable CPC when compared to PLGA-composites with different characteristics. However, the here presented results revealed that 10 % GDL degrades at a faster rate than PLGA microspheres generating earlier CPC porosity and therefore significant more bone formation within the material after only two weeks of implantation. Interestingly, 30 % GDL does lead to less bone formation than 10 % GDL possibly because, although bigger pores will be created, there is a smaller degree of interconnectivity at micro- and nanoporosity level which results in i) less fluid flow throughout the material which ends in slower GDL degradation rate and ii) this less interconnected porosity does not allow the necessary space for continuous bone ingrowth throughout the composite. In addition, a different initial structure, less granulated and more compact, of the CPC was observed under the SEM microscope between the 30% GDL and the other three composites. This different structure can also limit the ability of the bone to grow through the CPC. After 12 weeks implantation similar material degradation and bone formation occurred for the CPC-10% GDL and the CPC-PLGA samples. This reveals that though PLGA degradation occurs at slower ratio than GDL, when it occurs, composites can be degraded to a large extend and new bone tissue can be formed instead. Both composites exhibited a very good biocompatibility and replacement by bone tissue. Though the results indicated around 65 % of bone formation in the ROI when 10 % GDL or PLGA are used as porogens, it has to be taken into account that due to the trabecular structure of bone, there is an underestimation of new bone formation because a ROI filled with 100 % trabecular bone and no remainders of implant material leads to about 70 % bone tissue in the rabbit femoral condyle. Meaning that the obtained 66 % trabecular bone formations corresponds to approximately almost complete bone regeneration (94.3 %) of the ROI filled with bone tissue (including both bone trabeculae and interstices containing bone marrow or fat).

In previous investigations, adverse inflammatory response was found when gelatin microspheres where injected as a bone substitute in combination with injectable CPC after 4, 8 and 12 weeks implantation. Several reasons could contribute to this adverse effect of gelatin-CPC composites such as poor sterilization process or unwashed remainders of the cross linking agent glutaraldehyde (GA). The gelatin sterilization by ethylene oxide and the glycine washes performed to diminish unreacted GA prior to implantation in this study led into a good response of the bone tissue in the vicinity of the injected material Ethylene oxide was the sterilization method of choice due to the previously reported chain degradation and cross-linking effect of gamma radiation in gelatin-containing materials. Glycine washes were used as described before to reduce the unreacted GA which can be the cause adverse tissue reactions. The application of both combined treatments have successfully improve the CPC- gelatin bone interaction and new healthy bone formation can be seen in the implantation area.

Finally it needs to be emphasized that though GDL is a commonly used additive in food industry and it is applied in a variety of products such as milk, cheese or cookies among others, the use of GDL in biomedical applications has been also been investigated (i.e. alginate hydrogels for tissue engineering applications in which GDL is being used to control gelation rate, or as a part of cell-biomaterials constructs for regenerative medicine). After the here reported promising in vivo bone response, GDL emerges as a good CPC porogen candidate because of its fast degradability and subsequent rapid bone formation. Moreover CPC-GDL off the self availability where no complicated preparation protocols in order to fabricate 3D structures such as fibers or microspheres are required, the possibility of room temperature storage and low pricing in comparison to other materials commonly used as porogens such as PLGA or gelatin highlight GDL as a promising CPC porogen.

Conclusion

Injectable CPC with porogens of different nature (PLGA, gelatin and GDL) applied as bone substitute materials in femoral condyle bone defects in New Zealand White Rabbits demonstrated a significantly faster creation of CPC porosity, degradation and bone substitution for both CPC-GDL composites (CPC-10%GDL and CPC-30%GDL) compared to CPC-PLGA or CPC-gelatin after 2 weeks implantation. After 12 weeks implantation CPC- PLGA and CPC-10%GDL undergo complete degradation and replacement by bone tissue whereas significantly slower degradation occurs in CPC-30%GDL and CPC-gelatin

GDL emerges as a suitable CPC porogen when fast degradation and replacement of the material by newly formed bone is desired. Description of the Figures

Figure 1 : Liquid phase-dependent macroporosity in CPC-30%GDL samples.

Figure 2: 4 mm circular ROI used for the histomorphometrical measurements.

Figure 3: SEM micrograph of GDL powder.

Figure 4: SEM micrographs of CPC-PLGA (A) CPC-gelatin (B) CPC-10%GDL (C) CPC-30%GDL (D) composites.

Figure 5: Gluconic acid detection by RP-HPLC.

Figure 6: X-ray diffraction (XRD) analyses of the calcium phosphate cement discs. The XRD powder patterns of (A) CPC-GDL, CPC after 3 days incubation in PBS, CPC-GDL after 3 days incubation in PBS and GDL. (*= GDL peaks; += hydroxyapatite peak; #= β-TCP peak).

Figure 7: Histological sections at 2 weeks implantation of CPC-PLGA (A) CPC-gelatin (B) CPC-10%GDL (C) CPC-30%GDL (D) composites.

Figure 8: Histological sections at 12 weeks implantation of CPC-PLGA (A) CPC-gelatin (B) CPC-10%GDL (C) CPC-30%GDL (D) composites, * = visible non dissolved GDL crystals.

Figure 9: Quantitative evaluation of bone formation: Bone formation in the defect area (ROI) in the four composite types after 2 and 12 weeks of implantation.

(a) significantly different compared to CPC-10%GDL

(b) significantly different compared to CPC-30%GDL (c) significantly different compared to CPC-PLGA

(d) significantly different compared to CPC-gelatin

Figure 10: Bone formation relative to a ROI filled with trabecular bone.