- INTRODUCTION
Bone cements are widely used in orthopaedic surgery for the fixation of prosthetic devices, such as total hip or knee replacements to bone. Bone cement mainly consists of polymethylmethacrylate (PMMA). [4]
This polymer was reportedly first used in a medical application in the late 1930’s as a means to cover skull defects. Further hints of the excellent biocompatibility of this polymer in tissues came from World War II fighter pilots that had been injured during aircraft crashes. The first clinical use of PMMA mixtures (in 1938) was an attempt to close cranial defects in monkeys. In 1956 Judet and Judet were the first to introduce an arthroplastic surgical method. In 1958 Sir John Charnley was the first researcher who succeeded in anchoring femoral head prosthesis in femur with in site auto-polymerization of PMMA [11]. The cementing technique of that time involved mixing the liquid monomer and granular polymer components with a spatula in an open bowl. The surgeon would knead the resulting doughy mass, prior to finger packing and thumb pressurizing it into place. Subsequently the prosthesis was inserted and the bone cement was allowed to polymerize. Figure 1 shows (A) an X-ray of a cemented total hip arthroplasty and (B) a removed prosthesis with surrounding bone cement mantle.
This initial technique showed a survival rate of about 85% after 15 years of follow-up, which left room for improvements. Later generation cementing techniques mainly evolved from clues from post-mortem studies and laboratory models of the cement bone interface in earlier generation cementing techniques. Second generation techniques aimed at improving the contact of the bone cement with the bone. The introduction of porosity reducing measures marked the start of third generation cementing techniques. Other third generation developments in cementing techniques are the use of prosthesis positioning devices, such as centralizers, that ensure correct placement of the prosthesis and a reliable thickness of the bone cement mantle. With these new techniques cemented total hip arthroplasty currently shows survival rates of 85–90% at 15 years and 80–85% at 20 years [3].
Fig. 1. Bars represent 5 cm. (A) X-ray in a frontal view of a patient with a cemented total hip replacement. The arrow indicates a region where the bone cement has loosened from the bone. (B) Photograph of a similar cemented prosthesis after removal from the patient. The bone cement was still firmly attached to the prosthesis [9].
Today most of the bone cements that are used in dentistry and orthopaedic surgery are made of PMMA. In dentistry, pure PMMA is used in prosthetic applications and the compositions that contain PMMA are used as a filler material. In orthopaedic surgeries, bone cements are mainly used as filler for filling cavities resulting from post-operative bone loss and for anchoring the metallic implants securely into the bone medulla. The cement serves, in addition to fixation, as an interface between the metallic implant and natural bone. It also transfers and distributes static and cycling loads due to daily activities.
PMMA is used in many medical applications such as blood pumps and reservoirs, membranes for blood dialysers and in vitro diagnostics due to its excellent biocompatibility and ease of manipulation. Its optical properties also make it ideal for implantable ocular lenses and contact lenses and its physical and colouring properties make it a good denture material [9].
- ChemIcal ComposItIon of Bone Cement
Most of the currently available acrylic bone cements have similar compositions. Some variation occurs across commercially available cements [12]. Acrylic bone cements consist of two parts, namely powder and liquid. Powder part contains PMMA beads, an initiator such as benzoyl peroxide and BaSO4 or ZrO2. Liquid part is mostly composed of methylmethacrylate monomer (MMA) (Figure 2). Some commercial products include butylmethacrylate. An activator is also present in order to accelerate the polymerization reaction such as dimethyl-paratoluidine. Liquid component contains an inhibitor such as hydroquinone to prevent self polymerization of the monomer before use.
![Figure 1 [formula]](Bone-Cements_oktar_files/image003.gif)
Figure 2: Chemical formula of the PMMA monomer subunit, methylmethacrylate. [6] |
The composition of the liquid and powder parts of the acrylic bone cements are summarized below:
Methylmethacrylate (monomer) 85%
Comonomer: butylmethacrylate (binding agent) 15%
Dimethyl-paratoluidine (activator) 2–3%
Hydroquinone (stabilizer or inhibitor)
The powder:
Polymethylmethacrylate (PMMA) 90%
Benzoyl peroxide (initiator) 2–3%
BaSO4 or ZrO2 (contrast agent) 4–8%
General chemical compositions of commercially available bone cements are presented in Table 1.
Table 1. General chemical compositions of various commercially available bone cements [7]
2.1 Polymer / Monomer Ratio
Most of the commercial bone cement formulations have a polymer / monomer ratio of 2/1 (w/v). When this ratio increases (by increasing the amount of PMMA powder or by decreasing the amount of monomer) the viscosity of the cement dough increases which in turn results in the difficulty in workability and creates problems in applying dough deep in the bone. On the other hand smaller powder / liquid ratios give smaller peak concentrations of free radicals, larer peak temperatures and higher monomer release to the surrounding tissues [11].
2.2 Initiator
Benzoylperoxide (BPO) is the most common initiator in bone cement formulations and is added in the powder component. There are some other initiators such as tri-n-butylborane (TBB) used in the commercial product, Bonemiteâ. The chemical formulation of BPO and TBB are given below.
Chemical formula of BPO:
Chemical formula of TBB:
B(C4H9)3
After mixing the liquid and the powder parts, initiators produces radicals at room temperature and starts the polymerization reaction causing an increase in temperature.
2.3 Accelerator / Activator
Most of the bone cement formulations contain a tertiary aryl-amine, N,N-dimethyl-p-toluidine (DMPT), as an accelerator for polymerization reactions. DMPT is added in the liquid part and is water soluble to a small extent. It causes the decomposition of the BPO in a reduction-oxidation process by electron transfer that produces a benzoyl radical and benzoate.
2.4 Inhibitor
In order to prevent self-polymerization during storage, liquid components of bone cements contain an inhibitor as a radical scavenger. Hydroquinone is usually used for this purpose. It was shown that hydroquinone could be replaced by less toxic materials such as food grade di-tert-butyl-p-cresol.
3. BONE CEMENT PREPARATION
PMMA particles are ground by a water-cooled analytical mill (Tekmar, Janke&Kunkel GMBH Co. KG) and then sieved (using Endecotts Octagon 200 test sieve shaker, England) through 150 and 50 mm meshes for 20 min with a vibration frequency of 50 s-1. The average particle size and size distribution is obtained by a particle size analyzer (Malvern Mastersizer, Malvern Instruments Ltd, UK). The composition of commercially available CMW1 bone cement (Table 2) is chosen as a reference. Approximately similar percentages of the constituents with the same ratio of the solid/liquid components as in CMW1 are used in the preparation of compositions.
For the preparation of the solid part of the bone cement, weighed amounts of chemicals (PMMA, BPO and HA) are mixed intensively and kept in a container. The required amounts of chemicals (MMA and DMPT) are mixed in a glass bottle for the liquid part. Powder and liquid parts are freshly prepared, put on shelf and kept at room temperature of 23 ± 1 °C for 1 h prior to every experiment to achieve thermal equilibrium.
For the cement dough preparation, the powder and the liquid parts are manually mixed with a polyethylene spatula for 1–3 min and a homogeneous dough is obtained [11]
Polymerization of methylmethacrylate is an exothermic reaction that results in a doughy substance that self cures in a short time. The degree of polymerization is affected by the following: (1) the amount of accelerator and initiator in the powder and liquid monomer; (2) wetting caused by the monomer mixing with the powder; (3) the type of mixing used (e.g. hand mixing, centrifugation or vacuum mixing); (4) the pre-chilling of the monomer; and (5) the presence of oxygen.
The polymerizing process transforms the initial thick liquid to a soft deformable material and finally to a rapidly hardening cement with an associated increase in temperature due to the exothermic polymerization. It is critical that the polymerization process be reproducible so the surgeon can properly apply the cement.
Table 2 Chemical composition of CMW1 [11]
4. THERMAL cHANGES DURING CURING
The polymerization reaction releases a large amount of heat. A given amount of cement releases less heat if it is spread out into a thin layer. With a 4-mm layer of cement between the bone and the implant, the amount of heat released remains under the threshold at which coagulation of bone proteins or collagen occurs. To minimize the risk of bone necrosis, the cement mantle should be 2-4 mm thick [12]. Figure 3 shows the change in polymerization temperature of the bone cement with time. The dough obtained after mixing the powder and liquid parts of the cement starts to solidify and sets after a few minutes.
Figure 3 Change in polymerization temperature with time [11]
Setting Time: This is the time needed for the cement to harden once the two components are put in contact with each other. The cement is liquid at first then thickens to a doughy consistency. The dough phase ends when the cement no longer sticks to the fingers. Then, further polymerization solidifies the cement. The setting time varies with the powder/liquid ratio of the cement, the temperature of the cement, and the temperature of the operating room. Cold cement in a cold room takes longer to set. The setting time decreases by 2 min when the temperature increases from 20 to 27 °C. The operating room temperature must be kept constant to avoid modifying the setting time [12].
Dough Time: It is indicated by failure of the material to stick the surface of a surgical glove.
Working Time: It is the time interval between dough and setting time.
5. PropertIES of Bone Cements
Viscosity: The viscosity is one of the important parameters for bone cements. The dough should attain proper texture, give enough time to surgeon to use it with maximum workability and maximum penetration when pressed into the cancellous bone without getting harder. For good intrusion into the trabeculae, the cement should have the desired viscosity [11]. Many authors have suggested that low-viscosity cements should be preferred. A cement may take longer to harden than expected if the viscosity is too low, resulting in gaps, laminate formation or malalignment if the prosthesis moves before the cement fully hardens. On the other hand, a high viscosity cement may introduce cement–bone and cement–implant gaps because of poor cement penetration into the bone tissue and/or poor cement flow properties. Cement viscosity is increased by the addition of fibres, greater molecular weight of the polymer, solubility of the polymer in the monomer, variation in the powder composition or bead size distribution, and the temperature of the cement components [7].
Porosity: Cement is a porous material that contains a variable volume of air bubbles. Pore formation arrises from various sources: Air initially surrounding the liquid monomer or powder constituents, entrapment of air during wetting of the powder by the liquid monomer, entrapment of air during mixing of the constituents, the boiling or evaporation of the volatile liquid monomer during the curing stage and entrapment of air during the transfer of the dough to the syringe. Porosity varies across cements from 5% to 16%.
Volume changes: Polymerization changes the volume of the cement : the mixture shrinks at first, then expands during the heat release phase, and finally shrinks again when it cools. In theory, the monomer loses 20% of its volume. Given that the cement is composed of about one-third liquid and two-thirds powder, the loss of volume is about 8%. [12]
6. MECHANICAL PROPERTIES OF BONE CEMENT
Attempts to characterize the mechanical properties of PMMA have accompanied its application in orthopedic surgery. The majority of the objective data reported on the mechanical testing of this material have been published long after its initial orthopedic use. Despite the abundance of this data, little is known about the mechanical behavior of PMMA. In addition, variability in testing modalities and techniques has contributed to this problem and has led to inconsistent reports.
The materials commonly used in orthopedic applications must be put through the rigors of laboratory and clinical testing. With the recent emphasis on the higher standards imposed by the Food and Drug Administration, substantial data is now necessary to approve a material for new clinical use. As a result, a broader spectrum of mechanical testing has been used to more accurately characterize a new material. Application of this technology has led to the re-examination of PMMA in its current form. Furthermore, the jurisdiction of the American Society for Testing and Materials has suggested the need for the standardization of mechanical testing and the reassessment of previously published data on the mechanical properties of PMMA.
Mechanical strength of bone cement is very important since most of the aspectic loosening is related to fracture of the PMMA cements. For bone cements mechanical properties are effected by various factors and it is not easy to report strength characteristics of all new formulations because each differ from each other. Some of the factors effecting mechanical properties of bone cements are; composition of cement parts, weight-average molecular weight of the polymer part, porosity, type of sterilisation of the constituents and mixing methods.
The mechanical properties of PMMA have been reported by different authors (Figure 3). In 1984, Saha and Pal published a comprehensive review on the mechanical properties of bone cement [14]. They discussed the mechanical properties of PMMA and re-evaluated its static and dynamic properties.
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Figure 3: Graph representation of the reported value ranges of the static mechanical strengths (MPa) of four common PMMA bone cements. PMMA is strongest in compression (A); Simplex P have consistently performed better in flexural strength testing (B); data on the static shear strength of PMMA bone cements has been limited (C); and PMMA bone cement is weakest in tension (D). |
Static mechanical properties, such as compressive, tensile, and shear strengths are properties usually included in the mechanical characterization of any material. The study of compressive strength typically consists of an applied axial load to a cylindrical plug made of the material being tested. The applied load results in “strain” within the specimen and eventual material failure in compression. The load at which PMMA fails in compression is the ultimate compression strength (MPa). Determination of the ultimate tensile strength, or the ultimate shear strength of PMMA, also can be performed by subjecting the specimen to stresses in tension or shear until failure. The relationship between any of these stresses and the corresponding strain invoked within the material can be represented by the slope of the line within the linear portion of the stress-strain curve and is termed the modulus of elasticity. Generally, this value reflects the characteristic stiffness of the PMMA specimen.
Static tests report the failure of a material after a single load. Although large amounts of data have been accumulated from static compression, tensile, and shear mechanical testing, the prediction of PMMA in its clinical performance cannot be made from these results alone. Mechanical tests that mimic the mechanical milieu of PMMA in its clinical application are better suited for this task. A three- or four-point bending study is a static test that invokes compressive, tensile, and shear loads to PMMA specimen. The determination of the flexural strength as a result of this test model may be a more accurate static assessment of the clinical strength of PMMA. Static torsional testing of PMMA constructs should be considered as well.
The dynamic material properties of PMMA also have been evaluated. Di Maio et al have provided data that may better reflect the clinical behavior of PMMA cement. The dynamic mechanical evaluation of PMMA has included an abundance of fatigue testing. In this instance, PMMA failure is seen after fatiguing the specimen with frequent applications of submaximal loads.
The applied loads typically are less than loads associated with ultimate PMMA failure. The “cycling” of these smaller loads mimic the clinical loading of PMMA cement. In 1987, Davies et al. compared the mechanical strengths of commonly used commercial PMMA cements. Although no difference in the static tensile strengths of the tested cements was found, the fatigue properties of Simplex P were superior to other cements tested. These results suggest that although these materials may behave similarly during static testing, their differences in fatigue properties must be considered when considering their potential clinical performance.
It has been suggested that the failure of PMMA cement occurs as a consequence of “cracking” of the material. Crack initiation occurs from “voids” within the cement. These voids then act as stress risers, which result in cracks that propagate within the cement construct. Crack initiation and propagation eventually leads to the structural failure of this material. A reduction in the number and size of the voids within the cement would theoretically limit this process. Thus, the “porosity” of PMMA cement preparations has always been a major source of concern, and has proven to be one of the reasons for poor mechanical performance. Therefore, limiting the presence of voids has become an important issue in the quality control of the cement used clinically and can be used to judge whether a variation in cement preparation and handling is beneficial [6].
7. BONE CEMENT MIXING METHODS
Bone cements are usually a two-component system–a liquid monomer and a granular powder component. At the time of mixing, the components are usually hand mixed in a bowl. However, with the use of vacuum mixing or centrifugation after mixing, the cement porosity can be reduced (either in overall amount and/or
through the reduction of pores size, thus improving the mechanical properties of the cured cement).
In manual or hand mixing, the powder component is added to the liquid in a polymeric bowl. Then these components are stirred with a PP or PE spatula with a speed of 1 or 2 Hz for 45-120 seconds.
In centrifugation mixing the hand-mixed dough is immediately poured into a syringe that is than promptly placed in a centrifuge and spun with a speed of 2300-2400 rpm for 30-180 seconds.
In vacuum mixing methods generally 5 to 100 kPa of vacuum is applied for 15-150 seconds to the dough during mixing stage. In some of the vacuum mixing processes air within the powder is evacuated before mixing. Greater monomer evaporation may occur if the applied vacuum is too high during vacuum mixing.
Mixing of bone cement under vacuum generally has been reported to decrease the setting time, compared to cement mixed at ambient pressure [7].
8. ADVANTAGES AND DISADVANTAGES OF BONE CEMENTS
Poly(methyl methacrylate) (PMMA) is used in a wide variety of medical and dental applications and is currently the only material for anchoring cemented arthroplasties to the contiguous bones. In this application the main functions of the cement are to transfer body weight and service loads from the prosthesis to the bone, and increase the load carrying capacity of the prosthesis–bone cement–bone system.
The main advantages of cemented prosthesis lay in their excellent primary fixation, good load distribution between implant and bone, and the fact that the technique allows fast recovery of the patient. However, despite the relatively good success rate of implant fixation with acrylic-based bone cement, a number of persistent problems are encountered.
The disadvantages are related mainly to the necrosis associated to the use of the cement and the relatively poor mechanical behavior of the cement.
The major problem involved with the cement is the heat generated during the polymerization. The high local temperatures, reached due to the exothermic chain propagation reaction, have a deleterious effect on living cells and extracellular matrices. Such denaturation reactions lead to necrosis of adjacent tissues [1]. It is postulated that the cement has a role in thermal necrosis of bone, impaired local blood circulation and predisposition to membrane formation at the cement–bone interface. All these phenomena have been attributed to the high exothermic temperature of the cement, amounting to about 67–124°C, depending on the cement formulation. The temperature increase is governed by the amount of polymerizing monomer and the chemical composition of the cement components. Moreover, the chemical necrosis has also been postulated to be due to the release of unreacted monomer. Hence, it could be of clinical importance to reduce the residual monomer content, without influencing cement quality. On the other hand, either the shrinkage of cement during polymerization or the poor mechanical behavior of the cement contribute to the aseptic loosening of the implant [2].
9. ADDITIVES
9.1. Radiopacity:
It is essential that bone cements are radiopaque, as it is vital for radiological detection. However the radiopacity of orthopaedic bone cement itself, is very limited, due to the low density of the polymeric material which consist of molecular structures containing light elements such as hydrogen, oxygen and carbon. Radiopacity has been frequently achieved by addition of X-ray contrast materials, such as barium sulphate and zirconia, which are known to alter the biological and mechanical properties of the bone cement. These cements are known to be brittle in nature and the microstructure of the in situ polymerising cement is beset with inherent flaws. The predominant methods of rendering biomaterials radiopaque have been through the addition of heavy metal salts, however the major problem with this method, is the heterogeneity of the resultant matrix due to the incompatibility of the resin and the metal salt. Radiopaque polymer blends are generally used in bone cements, which are essentially a physical mixture forming a heterogeneous dispersion with the radiopacifying agent being distributed in the polymer matrix. Particulate barium sulphate incorporated in the matrix of bone cement, lowers the ultimate tensile strength and fracture toughness and reportedly has a greater effect on the mechanical properties in comparison to zirconia [13].
9.2. Antibiotics:
The concept of using the bone cement as a depot for antibiotics makes sense, as it allows delivery of antibiotics directly to the site of infection. Buchholz and Engelbrecht first reported on the possibilities of mixing antibiotics in bone cement in 1970. They considered gentamicin sulphate to be the antibiotic of choice because of its wide spectrum antimicrobial activity, its excellent water solubility, its thermal stability and its low allergenicity. Furthermore, gentamicin could still be found in tissues surrounding the implant for over years. This was thought to confer long-term protection against haematogenous infections. Gentamicin is a naturally occurring antibiotic produced by the bacterial strain Micromonospora purpurea and has been in clinical use now for over 50 years. It is a so-called aminoglycoside and this class of antibiotics has a concentration-dependent antibacterial activity. Apart from gentamicin, other antibiotics have also been used as an additive to bone cement. The combination of erythromycin and colistin is an example that made it to a commercial product, but this has two drawbacks compared with gentamicin. The antibacterial spectrum is narrower and it inhibits bacterial growth instead of killing bacteria. Adding gentamicin or other antibiotics to cement does not seem to alter the mechanical properties of bone cement. Several bench studies have compared the mechanical resistance of cements with and without antibiotics. These studies used fatigue tests, which are the best tests for simulating the long-term behavior of cement. Adding antibiotics in the amounts used in clinical practice caused no detectable alterations in the mechanical properties of the cement.
Prophylactic antibiotic therapy by the systemic route has been found effective in many studies. Consequently, the relative advantages and drawbacks of systemic and local prophylaxis must be discussed. The only comparative, randomized study compared the 2-year infection rate in 821 total hip replacements with gentamicin-loaded cement and 812 with standard cement and systemic prophylactic antibiotic therapy. Rates of deep infection were 0.4% and 1.6% in these two groups, respectively, and rates of superficial infection were 8.6% and 6%. These data suggest that gentamicin-loaded cement may decrease the risk of deep infection, which is serious, at the expense of a small increase in superficial infection, which is usually nonserious and easy to treat. Systemic prophylaxis is far more costly than gentamicin-loaded cement. No data are available on the efficacy of combining these two methods.
Anticancer agents such as methotrexate and cisplatin can be mixed into bone cement to provide a local curative effect after removal of a malignant tumor. Numerous experimental studies have confirmed that the anticancer agent diffuses in situ and retains its intrinsic properties. This method, which was developed by Professor Hernigou, is a simple means of providing adjuvant local chemotherapy. However, systemic chemotherapy and radiation therapy, when indicated, should also be used [12].
The release profile of antibiotics from bone cements follows a typical curve. A high initial release rate allows high initial concentrations, particularly in close contact with the bone cement. Subsequently, the release rate declines sharply, and a slow long-term release follows. Despite the fact that the release of antibiotics continues for a long time, for practical purposes it can be considered to be incomplete. After years of implantation it has been shown that over 80% of the antibiotic incorporated initially is still present isolated in the bone cement. Extended presence of antibiotics, as a result of the release characteristics described above, has not been shown to be as effective as the initial release.
In a recent study by Hendriks et al.(2003); increased release of gentamicin from acrylic bone cements under influence of low-frequency ultrasound was studied. They investigated the possible effect of ultrasound on antibiotic release from bone cements. Samples were made of three commercially available gentamicin-loaded bone cements. Part of the samples was allowed to release gentamicin for 3 weeks before insonation. An insonation device produced an ultrasound field at a frequency of 46.5 kHz. The samples were exposed to the ultrasound field or not exposed to it as a control. The amount of gentamicin released was measured by fluorescence polarization immunoassay. It was found that there is an increased release of gentamicin from fresh gentamicin- loaded bone cement under influence of low-intensity, low-frequency (46.5 kHz) ultrasound [8].
9.3. Fibers:
Although the use of poly methylmethacrylate (PMMA) as a self curing bone cement in fixation of prosthetic components in total-joint arthroplasty has received widespread acceptance, its use however, is not without complications. It has a weak and brittle structure and can consequently lead to failure of joint replacement due to its fracture. Therefore there is a need for an improvement in its mechanical characteristics [15].
One of the efforts for improving the properties of bone cement involves dispersing small quantities of enforcing materials such as; carbon, graphite, aramid, bone particle, polyethylene, titanium, ultra high molecular weight polyethylene, or PMMA fibers in cement matrix. Although most of the results regarding the properties of these filamentary composite materials have been encouraging, the biocompatibility issues regarding some of these fibers are not known. None of these reinforced cements have been approved by the FDA for clinical use [11].
10. MODIFICATIONS OF BONE CEMENT COMPOSITONS
Poly(methylmethacrylate) (PMMA) bone cement, used to fix implants into the bone, produces good surgical results if used correctly. However, prostheses do eventually become loose and the breakdown of the cement mantle is a factor in this failure. Limitations of PMMA cement, which lead to problems with the fixation of the implant, include its mechanical characteristics and its influence upon surrounding bone, associated with the polymerization reaction [16].
A bone cement with higher fracture toughness and higher creep resistance may be beneficial. One method for improving the properties of bone cements involves the development of new formulations.
10.1. Bioactive Bone Cements:
PMMA cement has widely been used in various prosthetic replacement operations. However, it has several inherent problems such as non-bone-bonding character, relatively low mechanical strength, and high heat generation during its polymerization [16].
Bone-PMMA bone cement interface is known as one of the weak-link zones in the prosthesis-bone cement-bone construct because it does not adhere to bone [17].
A good interface between the cement and bone is important and due to the high exotherm of the polymerization reaction and the toxicity of the MMA, bone necrosis can occur. The result of this is the production of a fibrous capsule around the implant, which allows micro-motion to occur, causing pain to the patient and a space for wear particles to accumulate. In attempts to overcome these problems, bioactive bone cement, which has a character of bonding directly with the living bone tissue in a few weeks, can be used [5]. Bioactive materials are able to react in a variety of ways with living tissue, rapidly forming chemical bonds. Bioactive materials are the ones that achieve a specific biological response at the interface of the biomaterial and form a bond between the tissues and the material. Generally the formed layer is biologically active hydroxycarbonate apatite, which is chemically and structurally equal to the mineral phase of the bone. Bioactive bone cement does not generate high heat when it hardens, and has a significantly greater mechanical strength than PMMA cement after hardening.
10.2. Calcium Phosphate Bone Cements:
Acrylic-based bone cements are widely used in dentistry and orthopaedic surgery. Intense research is being carried out to improve the thermal, mechanical and biological properties of bone cements. Some studies include addition of small quantities of ingredients such as carbon, graphite, aramid, bone particles, polyethylene, titanium, ultra high molecular weight polyethylene, PMMA fibers, tricalciumphosphate (TCP) or hydroxyapatite (HA) in the cement matrix.
HA is a biocompatible material and since it is osteoconductive it strongly integrates with bone. Therefore, addition of HA into bone cement formulations enhances biocompatibility and improves mechanical strength. HA is the inorganic material that forms the mineral phase of bone and its structure is calcium phosphate containing fluoride hydroxide chloride given by the general formula of Ca10 (PO4)6 (F,OH,Cl)2. Mechanical and thermal properties of the cement depends on the HA amount added. It was reported that HA addition, depending on the amount, size and surface properties of the particles, cause an increase or decrease in the mechanical properties of the bone cement. HA is reported to reduce water uptake of the cement affecting its mechanical properties. It is also reported that addition of HA into the cement formulation caused a decrease in curing temperature, but did not effect the curing time. However, the addition of excess amounts of HA prevents a smooth, pliable cement dough being obtained. The ideal case is, therefore, to achieve optimum increase in mechanical strength and decrease in curing temperature while retaining the workability of the cement dough. All these studies indicate that further research is essential on the mechanical and thermal properties of HA-impregnated bone cements.
In a recent article by Serbetci et al. [10] the thermal and mechanical properties of hydroxyapetite impregnated acrylic bone cements were investigated. In the study HA-containing acrylic bone cements were prepared and thermal and mechanical properties of the resultant cements were examined.
In order to achieve a proper and homogeneous distribution of HA particles in the polymer matrix, very low viscosity cement compositions were prepared by mixing poly(methylmethacrylate) particles with two different molecular weights. Addition of HA into the cement increased the viscosity while making workability easier. With this novel formulation, polymerization temperature decreased from 111 to 87 °C, compressive strength increased from 96 to 122 MPa and sufficient compressive fatigue strength of 77 MPa at 10 6 cycles was provided. HA-containing acrylic bone cements demonstrated higher mechanical strength than the reference cement. The screening tests by in vivo applications achieved biological compatibility. They concluded that HA-containing acrylic bone cement may effectively be used in the clinical field in the future.
11. CONCLUSIONS
Polymethylmethacrylate (PMMA) has been used almost 40 years as bone cements to fix orthopaedic implants to the bone and in dental applications. However, there are some well-known shortcomings in bone cement of this kind. The autopolymerization reaction of these cements is exothermic (i.e. producing heat), which can lead to necrosis of tissues. Furthermore, the surface of the polymerized bone cement is of a dense structure, which does not allow the bone to grow into the cement. The other shortcomings that have been reported include: poor adhesion to the bone, low mechanical properties, chemical necrosis of bone due to the release of unreacted monomer, and shrinkage during the polymerisation. Acrylic bone cement is a relatively brittle material, but reasonably strong in compression.
Several efforts have been made to improve the mechanical properties and biocompatibility of commercial bone cements by alteration of the mixing method of commercial bone cements, by the addition of fibers to the cements, and by the addition of mineral particles (e.g. hydroxyapatite, inorganic bone, or tricalcium phosphate) to the bone cement. Anticancer agents such as methotrexate and cisplatin, and antibiotics like gentamicin can also be mixed into bone cement to provide a local curative effect after removal of a malignant tumor and post-operative infections, respectively.
While developing new formulations of bone cement, one should take into account that alteration of one component can improve one property while causing a weakness in another property and all properties should be checked before any application.
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