Bone is a dynamic and highly vascularized tissue that continues to remodel throughout the lifetime of an individual. It plays an integral role in locomotion, ensures the skeleton has adequate load-bearing capacity, and acts as a protective casing for the delicate internal organs of the body.
Bone fracture is a substantial public health issue and the need for bone regeneration is increasing dramatically as the world population ages, Bone defects are one of the leading causes of morbidity and disability in elderly patients, leading to decrease in overall health and quality of life.
The high regenerative capacity of bone, particularly in younger people, means that most of fractures will heal well without the need for major intervention. Despite this, large bone defects, as observed after bone tumor resections and severe non-union fractures, lack the template for an orchestrated regeneration and require surgical intervention.
A treatment called grafting which includes either autografting or allografting, is used. Autografting (bone/tissue taken from the patient itself) is regarded as a gold standard, nearly eliminating the risk of reaction. While allografting (bone/tissue taken from another individual of the same species) is limited by high non-union rate with the host tissues.
Bone defect repair using the tissue engineering approach is perceived as a better approach because the repair process may proceed with the patient’s own tissue by the time the regeneration is complete.
Tissue engineering is the use of a combination of cells, engineering and materials methods, and suitable biochemical and physicochemical factors to improve or replace biological tissues. Tissue engineering involves the use of a tissue scaffold for the formation of new viable tissue for a medical purpose.
When skeletal defects are unable to heal on their own, bone tissue engineering (BTE), a developing field in orthopedics can combine material science, tissue engineering and regenerative medicine to facilitate bone repair.
Biomaterials for Bone Repair:
Materials scientists aim to engineer an ideal biomaterial that can mimic natural bone with cost-effective manufacturing techniques to provide a framework that offers support and biodegrades as new bone forms.
In the case of bone, materials should preferably be both osteoinductive (capable of promoting the differentiation of progenitor cells down an osteoblastic lineage), osteoconductive (support bone growth and encourage the ingrowth of surrounding bone), and capable of osseointegration (integrate into surrounding bone).
Many bone substitute materials intended to replace the need for autologous or allogeneic bone have been evaluated over the last two decades. In general, they consist of either bioactive ceramics, bioactive glasses, biological or synthetic polymers, and composites of these. The ideal basic premise, if following the tissue engineering paradigm, is that the materials will be resorbed and replaced over time by, and in tune with, the body’s own newly regenerated biological tissue.
Calcium Phosphate (CaP) Based Biomaterials in Bone Tissue Engineering:
Tissue engineering is promising to meet the increasing need for bone regeneration. Calcium Phosphate (CaP) biomaterials are of special interest as they share chemical/crystallographic similarities to inorganic components of bone.
CaPs have been used successfully in various drug delivery applications in the form of nanoscale (particulate system) to microscale (coatings) to macroscale (Calcium Phosphate cement (CPC), paste, scaffolds) for local delivery and in some cases for targeted delivery.
CaPs show excellent bioactivity, biocompatibility and biodegradability in the physiological conditions due to their chemical similarities to the inorganic part of bone and they are non-immunogenic.
CaPs are osteoconductive and osteointegrative. Cell adhesion is influenced by their surface roughness, percent of crystallinity, solubility, phase content, grain size, particle size, surface charge and surface energy.
CaPs have ability to promote cellular function, expression and form a direct uniquely strong interface with bone.
Resorption is dependent on the phase content of CaP, particle size, crystallinity and porosity.
Porosity of CaP is not important only for its mechanical and resorbability but also for ingrowth of bone. In porous form, CaP can permit the ingrowth of bone tissue and cell. Increasing porosity greatly enhances the surface area in contact with body fluids, thus leading to faster dissolution rate.
The wettability (or hydrophilicity) of CaPs is extremely important since surface energy is an important factor in osteogenesis regulation.
Categories of CaP:
Depending on nature, impurities and the presence of water, CaPs can exist in different phases. Exciting features of CaPs are their excellent bioactivity and biodegradability, but all CaPs do not have similar bioactivity and do not degrade at the same rate.
Bioactivity and degradation behaviour of CaPs generally depend on the Ca/P ratio, crystallinity and phase purity. CaPs are relatively insoluble at physiological pH 7.4, however they have increasingly high solubility in acidic environment i.e: below pH 6.5.
The most important property of CaP is probably its solubility in water. If the solubility of a Cap, e.g. HA, is less than the mineral part of bone, it degrades extremely slowly if at all. If the solubility of a CaP is greater than that of the mineral part of bone, it is degraded.
Two different categories of CaP should be distinguished: (i) CaP obtained by precipitation from an aqueous solution at or around room temperature (low- temperature CaP), and (ii) CaP obtained by a thermal reaction (high-temperature CaP).
Among all CaPs, the most acidic and soluble CaP is mono calcium phosphate monohydrate (MCPM). Hydroxyapatite (HA) and Beta-tri calcium phosphate (B-TCP) are most commonly used because of their osteogenic property and ability to form strong bonds with the host bone tissue. Solubility of B-TCP is much higher than the HA and thus B-TCP is termed as bioresorbable ceramic.
Development of bi-phasic calcium phosphate (BCP) based biomaterials consisting of HA and B-TCP are also used to control the degradation properties.
CaP as Scaffold:
Being a major constituent of bone, CaPs have been extensively studied as scaffold material for bone tissue engineering. Such biomaterials have a composition and structure very similar to the mineral portion of the bone tissue and have been considered appropriate to develop scaffolds.
CaPs are the most widely used bone substitutes due to their excellent biocompatibility. CaP scaffold would eventually degrade while the newly formed tissue takes over the time. The degradation rate of these scaffolds depends on the solubility of the type of CaP.
One of the characteristics required by bone tissue engineering for scaffold is that they must have porous and interconnected structure, enabling the migration and distribution of cells. Tri calcium phosphate is superior ceramic with respect to stem cell differentiation and osteoconduction in vivo. Porous ceramics are used as bone fillers for small defects, where osteoconduction is sufficient.
With the purpose to achieving the porous characteristic of bio ceramic (in this case CaP), paraffin spheres are used as porogens in the scaffold. Paraffin is choosen because it is an inorganic material with a low melting point, which make is possible to control the distribution and size of the pores, thus producing a macroporous and interconnected scaffold.
Based on findings, CaP scaffolds with macroporosity generated by paraffin spheres have interconnected pores and pore size compatible with the size of the spheres used. The scaffold enabled adhesion, proliferation and differentiation of the mesenchymal stem cells in osteogenic cells and therefore can be used as scaffolds for bone tissue engineering.
Nanostructured CaP Coatings:
Another important application of nanostructured CaP is coating on metallic or other implants to enhance the bioactivity and osteoconductivity of bioinert materials. CaP can facilitate osteointegration with natural bone through the formation of an apatite layer.
Calcium Phosphate as Bone Cement (CPC):
Calcium phosphate cements (CPCs) are used for bone loss and fracture as injectable materials to ?ll bone voids and to improve hardware ?xation in fracture surgery. CPCs are self-setting and form a biomaterial that is chemically similar to the mineral content of human bone, possessing biocompatible and osteoconductive properties. Some CPC compositions have shown a fast resorption rate, which may be bene?cial for the regrowth of new bone tissue, and some compositions have even been shown to stimulate bone tissue formation in vivo.
CPCs offer several advantages over conventional calcium phosphate bio-ceramics and acrylic bone cement that permit their use as bone graft and substitute, including moldability, minimum bone cavity, direct in vivo insertion, in situ setting, optimum bone implant contact, and biocompatibility and bioactivity. Another great advantage of CPCs is their lower temperature in vivo.
CPCs are limited in application to non or moderate load bearing musculoskeletal defects because they lack enhanced toughness, reduced brittleness and improved reliability i.e. the low mechanical strength. With the use of additives, we can overcome these issues.
Calcium phosphate bone cements have been shown to provide compressive strength of up to 80 MPa measured under application near conditions without a precompaction of the cement paste leading to lower porosity/higher strength
CPCs were discovered by Brown and Chow in the 1980’s. CPC can only have three different end products: apatite (PHA), brushite (DCPD), and amorphous calcium phosphate (ACP).
CPC are obtained by mixing one or more several reactive Calcium Phosphate powders with an aqueous solution to form a paste that hardens within a restricted period.
CPC consist of solid and liquid components. A combination of CaP with or without other Calcium compounds make up the solid components and the liquid component may be inorganic (sulfuric acid, phosphoric acid) or organic acids (lactic acid, tartaric acid)
Commercial CaP cements presently available are consisting of amorphous calcium phosphate and di-calcium phosphate dihydrate mixed with saline solution. Another is consisting of tetra-calcium phosphate and anhydrous di-calcium phosphate mixed with phosphate solution or water.
CPCs are different from traditionally used bone substitutes such as granules and blocks which are not in a paste form and do not sustain a rapid phase transition. Moreover, several substances such as antibiotics, antitumoral, or antin?ammatory drugs can be easily added to CPCs generating drug delivery systems.
Handling: Beside having excellent biocompatibility and bioactivity, two other main advantages are its injectability and self-setting capability in vivo at body temperature. However, without any physical, chemical and compositional modification, CPC normally possess a relatively long setting time, poor injectability and poor cohesion.
Setting time: The factors that promote faster setting kinetics can potentially reduce the setting time of CPC. Such factors are
Smaller particle size.
Higher setting temperature.
A low liquid to powder ratio (L/P ratio).
However, too short setting time may make CPC unworkable during total surgical
whereas longer setting time may cause inflammatory responses and disintegration of CPC
Cohesion and anti washout ability: Smaller particle size and control over the L/P ratio can be strategically used to strengthen particles interaction thus improving cohesion. Moreover, enhancing the viscosity of the mixing fluid by dissolving biopolymers such as sodium alginate, chitosan and modified starch, though can prolong setting time and hamper mechanical strength, has found to be effective in improving cohesion and anti-washout properties of CPCs.
Injectability: Injectability of CPC paste is the ability to stay homogenous during injection. Parameters such as decreasing particle size, increasing L/P ratio, adding ions or polymers and increasing the viscosity (best strategy) can be applied to improve injectability of CPCs.
Strengthening of CPC: To reduce brittleness of CPC and to improve their mechanical performance for load bearing applications, several research efforts are in progress which include, modification of cement with polymeric additives, reinforcement with resorbable as well as strong and tough fibres to the cement matrix or using dual setting cements.
Macroporosity: CPCs are intrinsically porous materials, with pores in the micro or nanometer range,20 but lack macroporosity, which is an essential feature for tissue colonization and angiogenesis. Two main routes have been explored to introduce macroporosity into CPCs by adding polymers: (a) foaming the liquid phase or the cement paste containing a polymer, and (b) loading the CPC with biodegradable polymers (e.g. microspheres (MSs) or fibers) that slowly degrade over time, resulting in a macroporous structure
The addition of citric acid enhances the mechanical strength of CPCs, and other studies have demonstrated that the use of chitosan and glucose with citric acid further improve the mechanical properties of CPC cements. Polysaccharides such as chitosan and sodium alginate have also been shown to enhance the washout resistance and mechanical properties of CPCs.
CPC Reinforced with Chemically Activated Carbon Fibers: The ?ber-reinforcement of ceramic materials enhances fracture resistance, but simultaneously reduces the strength of the composite. Combining strong C-?ber reinforcement with a hydroxyapatite to form a CPC with a chemical modi?cation of the ?ber surface allowed us to adjust the ?ber–matrix interface and consequently the fracture behaviour. Thus, we could demonstrate enhanced mechanical properties of CPC in terms of bending strength and work of fracture to a strain of 5%. CPC reinforced with chemically activated C-?bers is a promising bone replacement material for load-bearing applications.
Adding an organic phase to the calcium phosphate cement formulation is a very powerful strategy to enhance some of the properties of these materials. Adding some water-soluble biocompatible polymers in the calcium phosphate cement liquid or powder phase improves physicochemical and mechanical properties, setting time, macroporosity, injectability, cohesion and long-term degradation.
Incorporation of CPC with Chitosan: Chitosan is a linear polysaccharide composed of randomly distributed D-glucosamine and N-acetyl D-glucosamine units. Chitosan increases the flexural strength of a chitosan–CPC composite composed of TTCP–DCPA considerably, and the highest value was reached when 15–20 wt.% chitosan was incorporated into the CPC. In general, flexural strength decreases when the amount of chitosan increases ;20 wt.%. Similarly, compressive strength drastically decreases when chitosan increases to ;10 wt.%. Nevertheless, the compressive strength of CPC composites containing chitosan generally increases. An interesting property of chitosan is its ability to increase the anti washout resistance of CPC but not injectability.
Microporous CPC layer can also be incorporated with chitosan to enhance mechanical strength.
Incorporation of CPC with Alginate: Alginate is an anionic polysaccharide found in brown algae cell walls. The compressive strength of a CPC composed of DCPD–ACP containing sodium alginate increases as the concentration of polymer decreases. This was also observed for TTCP–DCPA and ?-TCP cements, in which the incorporation of low amounts of sodium alginate increases tensile strength.
Incorporation of CPC with synthetic polymers: Polymers are incorporated with CPC with the main advantage of increasing mechanical strength (decreasing ductility of CPC). They also avoid immunogenicity and disease transmission and possess flexibility in property controls.
Polyacrylic Acid: Polyacrylic acid (PAA) and its derivatives are capable of absorbing water many times their weight. Polyacrylates have considerable effects on mechanical properties. Compressive strength increases substantially to 55 MPa when ammonium PAA is incorporated into the CPC, which is contrasted with 25 MPa for a CPC without PAA. The increase in compressive strength can also be deduced from the reduction in composite porosity. Furthermore, adding PAA allowed the brittle CPC to become more ductile.
Polyesters and Polymethyl methacrylate (PMMA): Polyesters are thermoplastic polymers that contain an ester functional group in their main chain. Main functions of the polyesters in CPCs is to increase mechanical strength. However, these polymers are not water soluble, and therefore, they cannot be directly incorporated into the liquid phase of the CPC. PMMA is the most commercially important acrylic polymer. CPC can also be incorporated with PMMA to enhance its mechanical strength.
CPC for Drug Delivery:
In addition to acting as bone substitute, CPC can also be used for drug delivery for the treatment of different skeletal disorders such as bone tumors, osteoporosis, osteomyelitis. The highly microporous structure of CPC, after setting, allowing it to incorporate drugs into its structure.
The drug can be introduced either in the liquid or the solid phase of the CPC, but care must be taken that the physiochemical properties of the drug or protein do not change during the chemical reaction and setting of CPCs.
Different kinds of drugs including antibiotics, antimicrobial peptides have been incorporated into CPC for various applications.
Adverse Reactions of CPC:
Structure of CaP: Calcium Phosphate is a family of materials and minerals containing Calcium ions (Ca2+) together with inorganic phosphate anions. Some so-called calcium phosphates contain oxide and hydroxide as well. They are white solids of nutritious value
Monocalcium phosphate (MCP): It is an inorganic compound with the chemical formula Ca(H2PO4)2. (“ACMP” or “CMP-A” for anhydrous monocalcium phosphate). It is commonly found as the monohydrate (“MCP” or “MCP-M”), Ca(H2PO4)2·H2O. Both salts are colourless solids.
Dicalcium phosphate: It is the calcium phosphate with the formula CaHPO4 and its dihydrate. The “di” prefix in the common name arises because the formation of the HPO42– anion involves the removal of two protons from phosphoric acid, H3PO4. It is also known as dibasic calcium phosphate or calcium mono-hydrogen phosphate.
Tricalcium phosphate (TCP): It is a calcium salt of phosphoric acid with the chemical formula Ca3(PO4)2. It is also known as tribasic calcium phosphate. It is a white solid of low solubility. It exists as three crystalline polymorphs ?, ?’, and ?. The ? and ?’ states are stable at high temperatures. As mineral, it is found in Whitlockite.
Amorphous calcium phosphate (ACP or ATCP): It is a glassy precipitate of variable composition that is formed in double decomposition reactions involving a soluble phosphate and calcium salts (e.g. (NH4)2HPO4 + Ca(NO3)2) performed under carefully controlled pH conditions. The precipitate will either be amorphous tricalcium phosphate (ATCP) or calcium deficient hydroxyapatite (CDHA) (note CDHA is sometimes termed apatitic calcium triphosphate). The composition of amorphous calcium phosphate is CaxHy(PO4) z·nH2O, where n is between 3 and 4.5.
Hydroxyapatite (HA): It also called hydroxyl apatite (HA), is a naturally occurring mineral form of calcium apatite with the formula Ca5(PO4)3(OH). Pure hydroxyapatite powder is white. Naturally occurring apatites can, however, also have brown, yellow, or green colorations.
Polymethyl Methacrylate (PMMA) as Bone Cement:
Polymethyl methacrylate (PMMA) bone cement has been widely used for implant fixation in various orthopaedic and trauma surgery.
This cement is primarily composed of PMMA powder and methyl methacrylate (MMA) monomer liquid, with additions of hydroquinone to prevent premature polymerization, N, N-dimethyl-p-toluidine to accelerate the curing, barium sulphate or zirconium dioxide to add radiopacity, and dibenzoyl peroxide as an initiator.
PMMA is an acrylic powder that acts as space filler that creates a tight space which holds the implant against the bone and thus acts a grout.
PMMA has been used in surgical fixation of artificial joints for over 50 years. The primary function of bone cement is to transfer forces from bone to prosthesis.
The first clinical use of PMMA mixture was to close cranial defects in monkey in 1938. Surgeons used the heat stable polymer (Paladon 65) to close cranial defects in humans. The material was assembled in plates in laboratory and later moulded in the surgical suite.
The main advantage of PMMA is its rapid strength development that allows for fast recovery following joint replacement. Although the mechanical properties of PMMA may vary with temperature, environment, mixing procedure, porosity and strain rate, a typical range of values is as follows: elastic modulus 2.2-3.7 GPa, compressive strength 78-120 MPa and tensile strength 13.2-48.2 MPa. In vivo experiments have found that the mechanical properties increase during an initial period, generally days to month, followed by a slow decrease over the succeeding years. The fatigue behavior of PMMA bone cement in vivo depends mostly on the behavior of crack propagation.
The bonding at the cement-prosthesis interface depends on the surface roughness of the prosthesis. The strength of the interfacial bonding also depends on the type of the metal used to form the prosthesis. The bonding at the cement-bone interface relies upon mechanical interlocking from the protrusion of the cement into the trabeculae of the bone.
The main disadvantage of PMMA bone cement for use in joint replacement surgery is that approximately 10% of the patients may require a revision within 10 years.
Aseptic loosening often occurs which has been attributed to: lack of chemical bonding, mechanical failure of the cement, fibrous tissue formation resulting from heat induced necrosis. Another concern in the use of PMMA is suspected toxicity of the monomer.
Short term properties of PMMA show that bone cement is weak in tension and strong in compression. Long term properties of PMMA (creep, stress relaxation and fatigue) are the properties that can significantly affect the transmission of load into bone over the expected life or orthopaedic.
The overall mechanical properties of bone cement can load to long term stability of the replacement joints.
The production of wear particles from roughened metallic surfaces and from the PMMA cement promotes local inflammatory activity, resulting in chronic complication to hip replacements.
Bone cement generates heat as it cures and contracts and later expands due to water absorption. It is neither osteoinductive nor osteoconductive and does not remodel.
Bone conduction is depended on the conditions for bone repair as well as the biomaterial used and its reactions more than 60% of weight of bioactive ceramic powder should be added to PMMA powders to achieve satisfactory osteoconductive properties after setting.
PMMA along with various additives, gives the mixture a set of physical and chemical properties.
A new bioactive bone cement consisting of PMMA as an organic mixture and bioactive glass beads as an inorganic filler has been developed.
Structure of PMMA:
Chemical formula : (C5O2H8)n
PMMA is a vinyl polymer, made by free radical vinyl polymerization from the monomer methyl methacrylate.
Commercially available bone cement formulations are based on two components, a solid and a liquid phase, which are packaged separately and are mixed in situ in the operating theatre, prior to use. An important parameter of acrylic bone cement formulations is the solid: liquid ratio, which is 2: 1 in most of the cements. In some formulations, however, the solid: liquid ratio varies and may range from 2: 1 to 2.7: 1.
Solid Phase: The solid phase of the early cements mainly comprised spherical particles (beads) of poly (methyl methacrylate) (PMMA); however, currently marketed cements may contain low proportions of acrylic copolymers containing ethyl acrylate or methyl acrylate or even methyl methacrylate (MMA)–styrene copolymers in certain trademark cements. The prepolymer is present at a concentration of approximately 80 wt.% and it is prepared by suspension polymerisation. In this process, the monomer containing the radical initiator, e.g. benzoyl peroxide (BPO), is dispersed in an aqueous medium in which it forms droplets, which are stabilised by a surfactant agent under mechanical agitation and at a reaction temperature around 60–80°C. The solid phase also contains the initiator, BPO, in a concentration ranging from 0.75 to 2.5 wt.%. The initiator can be incorporated within the polymer particles or can be physically mixed. Another component of this phase is the radiopaque agent, namely barium sulphate or zirconium dioxide, that is incorporated as an X-ray contrast agent to allow X-ray follow-up of the prosthesis. The concentration of the radiopaque agent is generally around 10 wt.%
Liquid phase: The liquid phase consists of MMA monomer at a concentration of 95 wt.% for most of the formulations, although a few formulations contain small fractions of butyl methacrylate (BMA). It is customary to incorporate N, N-dimethyl-4-toluidine (DMT) as the activator, in a concentration range of 0.89–2.7 wt.%
Adverse Reactions of PMMA: The monomer is toxic and there is a potential for allergic reactions to cement constituents. The most frequent adverse reactions reported with acrylic bone cements are:
Transitory fall in blood pressure.
Loosening or displacement of the prosthesis.
Super?cial or deep wound infection.
Short-term cardiac conduction irregularities.