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BONE PLATES

by Burcu KAYIN

1.INTRODUCTION

Bone plates are surgical tools, which are used to assist in the healing of broken and fractured bones. The breaks are first set and then held in place using bone plates in situations where casts can not be applied to the injured area. Bone plates are often applied to fractures occurring to facial areas such the nose, jaw or eye sockets as seen in Figure 1 and Figure 2. Repairs like this fall into an area of medicine known as osteosynthesis (1-2-3).

Figure 1. Conventional bone plate used to repair jaw fracture (4)

Figure 2.Example of how a badly fractured face can be reconstructed using bone plates (4)

Plate osteosynthesis is an important technique in the treatment of femoral shaft fractures, particularly where an intramedullary nail may not be ideal, as in adult and childhood polytrauma, especially with head injury, pulmonary compromise, complex metaphyseal–diaphyseal or periarticular fractures, open fractures with vascular injury, and an excessively narrow intramedullary canal.

Open reduction and internal fixation with plates and screws has become a standard method of treatment for many types of fractures. However, the extensive operative exposure required to achieve anatomical reduction often results in devitalisation of the bone and surrounding tissues as well as evacuation of the fracture haematoma, which has osteogenic potential. While most fractures heal without complications, problems with delayed or non-union, infection and stiffness of adjacent joints are not uncommon. Therefore, more biological solutions have been sought (5).

 

2. The Properties of bone plates

The treatment of bone fractures aims at early and complete restoration of limb function. The fractured bone has loss its support function due to discontiniuity. Internal fixation restores continuity by splitting and or compression of the bone fragments. Immediate restoration of the mobility of the articulations avoids dystrophy which is one of the sequelae of therapeutical immobilization of the articulations and muscles.

Materials that are implanted for internal fixation should be strong, ductile, wear and fatigue resistant, have adequate Young's modulus, undergo no relaxation, maintain fixation, i.e. maintain mechanical properties during the healing period, be chemically stable for a given period, should not cause allergy and should not affect the infection resistance of the tissue or the organism. These materials should be safe and easy to apply, i.e. by conventional methods of application; they must be ductile enough to allow adaptation to the curved and twisted bone surface while maintaining strength. Bone plates are often used to support fractured bones. The bone plates are affixed using screws on to the bone over the fracture so that load is transferred via this bone plate while the bone is healing after which the plate is removed from the body. The bone plate should be biocompatible but should also have the appropriate mechanical properties as stated previously (6).

3. Metals for Implantation

Implantable materials or biomaterials are utilized to repair, assist or replace living tissue or organs that are functioning below an acceptable level. A wide range of metals and their alloys, polymers, ceramics and composites are used in surgically implanted medical devices and prostheses and dental materials. Most implanted devices are constructed of more than one kind of materials. Since the early 1900s, metal alloys have been developed for these applications.

Bone plates and screws are used in surgical procedures to stabilize or realign broken bones. The material and material processing techniques used in the construction of these products must be investigated. The material selection and design, product processing and product function is investigated by many researchers. There are a variety of different materials used for the construction of bone plates and screws. These are

• Stainless steel
• Titanium alloy
• Cobalt chrome
• Shape memory materials

The first metal alloy developed specifically for human use was "vanadium steel" in the early 1900's.The earliest successful implants were bone plates, introduced in the early 1900s to stabilize bone fractures and accelerate their healing. As early as the first bone plate implants, surgeons identified material and design problems that resulted in premature loss of implant function, as evidenced by mechanical failure, corrosion, and poor biocompatibility. Design, material selection, and biocompatibility remain the three critical issues in today's biomedical implants and devices.

These products, as seen in Figure 3, are used for the femoral head, bone screw, finger/toe plate and ankle plate. The appropriateness of the materials and processing techniques used should be investigated.

Condylar Plate 4.5mm	SHS Barel Plate with SCP Hole 38mm Barel Length
T Plate 4.5mm	SHS Barel Plate with SCP Hole 25mm Barel Length
T Buttress Plate 4.5mm	SCS Barel Plate with SCP Hole 95 0
L buttress Plate	SHS/SCS Lag Screw Top screw
Small T Plate 3.5mm	Small Oblique T Plate 3.5mm

 

Figure 3.   A multiplicity of plate forms and screw (7)

In 1920 attempts to employ surgical implants were still hampered by the limitation of available materials. In 1924 Zierald published the study on the reaction of tissues to a variety of metals. Iron and steel, the most widely employed materials at the time, were noted to dissolve rapidly and to provoke erosion of adjacent bone. Substantial discoloration of tissues was observed around specimens of copper, nickel, embedded in bones. A problem existed, as the metals, which did not produce discoloration, e.g. gold, silver, or pure aluminum, were all too soft or weak for most applications.

In 1926 18% chromium, 8% nickel stainless steel was introduced into surgical applications. This material was noted to be much more corrosion resistant in body fluids. This was stronger and more resistant to corrosion than the vanadium steel initially introduced by Sherman for his fracture fixation plates. Later in 1926, 18-8SMo stainless steel, which contained a small percentage of molybdenum, to improve the corrosion resistance in salt water, was introduced. This alloy became known as 316 stainless steel.

• Stainless steel

The austenitic stainless steels, especially Types 316 and 316L, are most widely used for implant fabrication. Stainless steel that has a low content of impurities and a passivated finish is entirely suitable for implantation in the human body. Forged stainless steel has a greater yield strength than cast stainless steels, but has a lower fatigue strength than other implant alloys. However, stainless steel is more ductile and more easily machined, and recent advancements have significantly enhanced its properties. Because a femoral component fracture with early designs, stainless steel is no longer used routinely, from the standpoint of erosion, biocompatibility, and fatigue life, stainless steel is inferior to other super alloys.

The only difference in composition between 316L and 316 stainless steel is the content of carbon. A wide range of properties exists depending on the heat treatment (annealing to obtain softer materials) or cold working (for greater strength and hardness). Even the 316L stainless steels may corrode inside the body under certain circumstances in a highly stressed and oxygen depleted region, such as contact under screws or fracture plates. Thus, stainless steels are suitable to use only in temporary implant devices, such as fractures plates, screws and hip nails.

• Titanium alloy

The next alloy to be introduced into orthopedic practice was titanium and its alloys. In 1947 possible applications for titanium surgical implants were considered. The pure metal had shown excellent inertness in an environment of seawater, so that corrosion resistance seemed likely to be good in the human environment. A few surgical implants were made and inserted into human subjects.

Attempts to use titanium for implant fabrication dates to the late 1930's. It was found that titanium was tolerated in cat femurs, as was stainless steel and vitalium (a CoCrMo alloy). Titanium's lightness and good mechanical and chemical properties are salient features for implant applications. One titanium alloy (Ti6Al4V) is widely used to manufacture implants. The main alloying elements of the alloy are aluminium (5.5 - 6.5%) and vanadium (3.5 - 4.5%). While the strength of the titanium alloys varies from lower than to equal to that of 316 stainless steel, when compared by specific strength (strength per density), the titanium alloys outperform any other implant material. Titanium nevertheless, has poor shear strength, making it less desirable for bone screws, plates and similar applications.

Titanium also tends to seize when in sliding contact with itself or other metal. Titanium-based alloys that have a high coefficient of friction can cause problems. Wear particles are formed in a piece of bone if a piece of bone rubs against the implant, or if two parts of an implant rub against one another. Therefore, implants of titanium upon titanium generally are not used as joint surfaces. Titanium derives its corrosion resistance to the formation of a surface oxide film. Under 'in vivo' conditions, the oxide is the only stable reaction product. 

Currently osteotemy equipment is made primarily of titanium and stainless steel. The broken bones are first surgically reset into their proper position. Then a plate is screwed onto the broken bones to hold them in place, while the bone heals back together. This method has been proven both successful and useful in treating all manner of breaks; however there are still some drawbacks. After initially placing the plate on the break or fracture the bones are compressed together and held under some slight pressure, which helps to speed up the healing process of the bone. Unfortunately, after only a couple of days the tension provided by the steel plate is lost and the break or fracture is no longer under compression, slowing the healing process.

• Shape memory materials

Bone plates can also be fabricated using shape memory alloys, in particular nickel titanium. Using a bone plate made out of NiTi surgeons follow the same procedure as is used with conventional bone plates. The NiTi plates are first cooled to well below their transformation temperature, and then they are placed on the set break just like titanium plates. However, when the body heats the plate up to body temperature the NiTi attempts to contract applying sustained pressure on the break or fracture for far longer than stainless steel or titanium. This steady pressure assists the healing process and reduces recovery time. There are still some problems to consider before NiTi bone plates will become commonplace. Designing plates to apply the appropriate amount of pressure to breaks and fractures is the most important difficulty, which must be overcome.

4. Corrosion of Implants

Any time a foreign material is placed inside the human body, the manner in which that material will affect the body must be considered. There are many causes that contribute to the corrosion of metals when placed inside the human body. After surgery the pH surrounding the implant is reduced to a pH between 5.3-5.6 typically due to the trauma of surgery. Infectious microorganisms and crevices formed between components can reduce oxygen concentration, both of which contribute to the corrosion of the implant (1-2-3).

Materials used must meet the biocompatibility constraints set forth by the Food and Drug Administration (FDA). If a company wants to introduce a material inside the body, the firm must demonstrate, through testing and analysis, the material's biocompatibility. Biocompatibility is the extent to which a material is compatible or friendly with the body. The major materials used for implants today include titanium alloy, cobalt chromium, and stainless steel. All these materials are considered biocompatible by the FDA. However, they all corrode and can cause complications inside the body.

A concern has recently surfaced in the biomedical field about the possible propensity for cobalt-chromium to cause cancer. Cobalt-chromium consists of the elements cobalt, chromium, nickel and molybdenum. There is a concern that the corrosion of cobalt-chrome in the wet, salty surroundings of the human body may be sending toxins streaming into the body, possibly causing cancerous tumors. Even though only about fifteen tumors have ever been reported at the site of an implant, many more could exist and go unreported (partially due to the age of most patients). Although these concerns have met some strong opposition in the industry, many companies are pushing towards safer materials. Such materials include titanium, inert fiber-reinforced composites, and ceramics. Studies involving titanium have illustrated that this material is generally well tolerated in the body.

Some other effects of corrosion exist in the implant materials. Skin conditions such as dermatitis have been reported from exposure to nickel. Cobalt shows signs of causing anemia by inhibiting iron from being absorbed into the blood stream. Ulcers and central nervous system disturbances have been detected as a result of chromium. Aluminum present in some implant materials may cause epileptic effects and Alzheimer's disease. Most of these side effects were the result of testing done outside the body in a different state from the implant. However, they do illustrate the possible hazards associated with the corrosion of implant materials inside the body. Because all metals corrode, preventing corrosion is difficult. The only apparent solution to this problem rests with reducing the amount of corrosion by choosing better quality materials. Efforts should also be made to use materials whose corrosion does not create adverse effects inside the body, such as titanium. These efforts can reduce the complications associated with corrosion. The common denominator of these types of complications is that they all increase pain and reduce the functional capacity of the implant. This leads to a subsequent loss in quality of life of the patient. Thus, the complications prevent total hip replacement from achieving its goal.

Orthopedic applications of metal alloys include arthroplasty, osteosynthesis and in spinal and maxillofacial devices. In order for these materials to perform successfully, they must have physical properties that allow the material to perform the function for which it was implanted, and the material must be biocompatible. In order for a material to be biocompatible, it must not adversely affect the physiological environment and the environment should not have detrimental affects on the material.

5. Composite bone plates

Metal bone plates made of stainless steel, Co–Cr and titanium alloys commonly used to treat bone fractures. However, the metal bone plates have very high elastic moduli (110–220 GPa) compared to that of human bone (17–24 GPa). The high modulus of metal bone plates results in almost all the load being taken by the bone plate and very little by the bone. Clinical research has shown that this is not desirable, as bone does not strain. The researchers have established that callus formation, ossification and bone union are hampered by the lack of strain in bone. This result in not only the fractured part but also the whole bone structure becomes osteoporotic. Bone plate should be just strong enough to promote the healing of fracture yet not so stiff as to hinder the bone union. Therefore, new types of bone plates are required which will have stiffness close to that of bone, yet biocompatible. Bone plates made of composites are possible candidates as they can be manufactured to stiffness similar to that of bone; they have high strength (or stiffness) to weight ratio, and non-corrosive. Further, the resulting composite behaviour can be tailored to nearly any requirement by choosing suitable matrix and reinforcement materials. It has been also shown that composite design possesses good static and fatigue resistance as against laminated or random design.

Fujihara (8) has reported that composites made of braided carbon fibres and epoxy resins have better mechanical properties than composites made of short or laminated unidirectional fibres. He has also demonstrated that braided fabric reinforced composites made of carbon fibre and epoxy resin could be used for bone plate application. It was also reported that carbon-epoxy is a biocompatible material.

Bone plates made of braided composites, because of their low stiffness, is subjected to considerable strain when subjected to loading. If the magnitude of strain is large, it is possible that matrix and fibre could experience relative displacement leading to fracture. Such a phenomenon is not desirable because of complexities involved in the removal of bone plate from human body. Therefore, it is important to calculate magnitudes of strain due to body loads to assess whether braided composites can sustain physiological loads or not.

Three specimens of the carbon–epoxy braided composite plates are tested in tensile loading. Tensile tests were conducted under a stroke of 3.0 mm/min by using an INSTRON 4206 universal testing machine at room temperature. The longitudinal and lateral strains are measured during the test. The average length, width and thickness of specimens are 120, 15.27 and 1.015 mm respectively. The stress–strain relation is linear up to failure and very little hardening/softening is observed and failure is sudden. The average strain at failure is about 1.5%. The average Poisson's ratio is 0.965, suggesting the material being not isotropic elastic. The composite material has average uniaxial tensile strength of 1.059 GPa and elastic modulus of 70.2 GPa. From the above description, the stress–strain relation of carbon–epoxy can be simulated using orthotropic elastic and a perfectly plastic model with a failure criterion similar to that of von Mises.

Bending behaviour of bone plates is the most important mechanical property, and is generally evaluated by bending moment and corresponding bending angle. The elastic bending moment capacity of only un-fractured bone up to the onset of plastic deformation (i.e. when stress in outer layer reaches yield value) is about 320 Nm.¹ The applied bending moment of 320 Nm induced about 0.5% strains in callus and 0.9% in composite plate. The magnitude 0.5% strain in callus depends on stiffness of callus, and its geometry. The value reported here gives a typically possible value. More clinical research is required to see whether or not 0.5% strain in callus is sufficient to stimulate callus growth. From the composite plate point of view, 0.9% strains can clearly be sustained by carbon–epoxy and it is very unlikely that de-bonding of matrix and fibre would take place. It should be mentioned that 0.9% strain developed when 320 Nm, which is moment carrying of un-fractured bone, is applied. In reality, moment applied will be less than this value, and therefore strains in composite will be less than 0.9%.

For a given loading, the longitudinal strain in bone and the bone plate are higher with composite plate than that with steel plate. This higher strain in bone could help growth of callus. The strains in composite bone plate are about 1%, when a moment equal to the ultimate capacity of un-fractured bone is applied. The results from stress–strain curves of the composite indicate that at 1% strain, there is no danger of fibre and matrix separation. These preliminary investigations suggest that braided composites may be used for bone plate design and manufacture.

6. Bioabsorbable Composite Materials

The metallic fixation devices, like screws and plates, have in some cases had disadvantageous side-effects. The modulus of elasticity of typical metals used in osteosynthetic devices is about 5 to 10 times that of bone. The differences in modulus of elasticity between bone and metals has been proved to cause the increasing of bone porosity and thinning of the cortical bone facing the plate during healing of the fracture (9). The corrosion of metals during implantation may in some cases cause strong reactions in the tissue. For these reasons, it is recommend that metallic osteosynthetic devices should be removed when the fracture has healed. The second operation increases human suffering and the cost of the treatment. Stress shielding is thought to interfere with blood circulation and the weakened bone may fracture again after implant removal. Thus, plates which exert less stress protection on bone and more flexible fixation of fractures are of increasing interest.

Because of the disadvantages of metallic fixation the use of bioabsorbable devices in bone fracture fixation was suggested already in 1960’s. Such devices retain their strength several weeks or months in vivo, support the healing fracture and are finally metabolised after fracture healing.

The advantages of bioabsorbable implants in bone surgery are significant: there is no need for removal operation, and osteoporosis associated with rigid metallic implants can be avoided or at least reduced and the bone itself heals better. The avoidance of removal procedures leads to financial benefits, psychological advantages, and it increases operative capacity.

During the last 20 years increasing interest has been concentrated in developing totally absorbable bone fracture fixation devices with properties resembling those of bone itself. Early experimental studies did not lead to clinical applications, evidently because of only modest mechanical strength properties implants (6).  

Regardless of the material or structure of the bioabsorbable tissue fixation implant, it must fulfill certain medical and physical demands to be safe for clinical applications. The medical demands are mainly bound with the biocompatibility of the material and its degradation products.

The following physical properties are essential for safe bioabsorbable implants:

(a) high initial strength;

(b) appropriate initial modulus;

(c) controlled strength and modulus retention in vivo.

The high strength is essential because the implant must resist mechanical stresses during a surgical procedure and it must carry external and physiological loads during the early stage of healing when the healing tissue is still weak. The appropriate modulus means that the material must not be too stiff or too flexible for the special purpose where it is used. The material should posses ductile behaviour so that it does not fracture with a brittle mechanism. The loss of strength and modulus in vivo must be in harmony with the increase of strength and modulus of the healing tissue. Non-reinforced bioabsorbable fixation devices have only modest mechanical strength. Therefore, such devices can be applied only in the treatment of non-loaded or slightly loaded fractures. For the safe treatment of loaded bone fractures (like femoral head and neck fractures) bioabsorbable implants must be reinforced to increase their strength.

Biodegradable plates suffer from warping, hollowing or substantial erosion inherent with the process of degradation. In order to solve the problem of insufficient mechanical strength, polymers with high crystallinity have been explored (10).

The research groups of Tormala and Rokkanen started in the late seventies the development of ultra-high strength, bioabsorbable composite materials to eliminate the application problems associated with the inadequate strength of non-reinforced bioabsorbable fracture fixation implants. In a series of materials, experimental and clinical studies the so-called self-reinforced, ultra high strength, bioabsorbable composites were developed from partially crystalline bioabsorbable polymers, like polyglycolides, polylactides and glycolide/lactide copolymers. These materials have high initial strength, appropriate modulus and strength retention time from 4 weeks up to 1 year in-vivo, depending on the material properties and implant geometry (6).

In the study of V.Hasırcı et.al. (10) it was mentioned that reinforcing elements such as fibers of crystalline polymers, fibers of carbon in polymeric resins, and particulate fillers, e.g., hydroxyapatite (HAP), have also been used to improve the dimensional stability and mechanical properties of biodegradable implants . The creation of interpenetrating networks (IPN) in biodegradable implants has recently been demonstrated as a means to improve implant mechanical strength . In order to improve the mechanical properties of IPN-reinforced biodegradable implants, biodegradable bone plates were prepared as semiinterpenetrating networks (SIPN) of crosslinked polypropylene fumarate (PPF) within a host matrix of poly(lactide-co-glycolide) 85 : 15 (PLGA) or poly(l-lactide-co-d,l-lactide) 70 : 30 (PLA) using different crosslinking agents. N-vinylpyrrolidone (NVP), 2-hydroxyethyl methacrylate (HEMA) were used as the hydrophilic crosslinking agents whereas ethylene glycol dimethacrylate (EGDMA) and methyl methacrylate (MMA) were used as the hydrophobic crosslinking agent. NVP was the control crosslinking agent and was partially or completely replaced by the other monomers. To determine the optimum crosslinking agent, tests were conducted comparing the amounts of crosslinking agent lost with the changes in mechanical properties and dimensional stability after in vitro treatment.

Bone plates were then prepared by replacing NVP with the optimum crosslinking agent. The mechanical properties were evaluated before and after in vitro treatment. In addition, HA was added to the composition to provide strength. Plates were also prepared by replacing PLGA with high molecular weight PLA. Bone plates prepared from PLGA and PLA were used as controls temperature.

The main disadvantage of current biodegradable materials is the premature loss of mechanical properties before the healing process is complete. In addition PLGA undergoes an autocatalytic degradation process which results in an accelerated degradation that leads to hollowing of the implant and its catastrophic failure.

A biodegradable polymer studied by S.Coşkun et.al.,(11) poly (3-hydroxybutyrate-co-3-hydroxyvalerate), appears to be a promising alternative. Microencapsulation of active substances have been utilized to protect drugs from inactivation and to achieve prolonged release. In the study osteoinductive substances obtained from bone marrow were planned for use in microencapsulated form to be implanted together with a biodegradable bone plate to achieve effective bone healing.

As a result of the study it was observed that an increase in the valerate content resulted in a decrease in the mechanical properties such as flexural strength and stiffness. Increase in Hap (hydroxyapatite) content improved mechanical properties almost in all categories. Bone plates with low valerate contents and 15% HAp had better mechanical properties.

In order to be able to adequately examine the implant in vitro degradation tests followed by mechanical testing must be carried out after incubation in various physiological media.

Implants with various geometries like plates have been developed from these materials and used succesfully in animal and in clinical studies. The main advantages of these implants are elimination of the need of the implant removal operation and biomechanically natural stiffness of the fixation leading to the rapid consolidation of the fracture.

7. Fixation technique

Plates find applications in many metaphyseal ares and are occasionally used to fix defects of long bones, typically in the radius, femur,tibia or mandible of the dog. For the case of the femur, the lateral aspect of the thigh is exposed by seperation of the gluteus maximus and the anterior fascia lata and by splitting or elevating the lateral musculature. The periosteum may be stripped or simply nicked in the areas where screw holes are placed. Removal of the periosteum reduces new bone growth on the periosteal surface and may retard bone regeneration under the plate.

Depending on the experimental objective, the plate fixation may or may not be accompanied by an osteotomy and creation of a gap in the bone. A  transverse osteotomy followed by compression of the bone ends is usually performed to gauge the effect of the biomaterial on the fracture healing response. Load-bearing gap filler materials can be evaluated by placing them in a plate-fixed lomg bone defect and compressing them using a self-compressing plate. Caution is required when compressing brittle ceramic-type defect fillers because thay may be subjected to high loads near their region of failure.

A commonly accepted fixation technique involves the use of the self-compressing or dynamic compression plate, as seen in Figure 4. This plate  has been designed with oval slots and spherical head screws. If the screw is located at one edge of the slot, it displaces the plate and away from the screw body when the head of the screw comes in contact with the side of the slot. This creates a compressive effect at the osteotomy site (12).

Briefly, the bone is located on the tensile side of the bone under physiological bending, which corresponds to the lateral side in the tibia and femur.A drill guide locates the hole off-set in the slot, and screw lengths are selected so that the screws span both cortices. The additional holes can be drilled and tapped, and the plate can be screwed in place. Following fixation of the plate, an osteotomy is amde in the bone, if healing is the phenomenon to be studied. The plate, already fixed in place,maintains allignment between the cut femoral components. If the plate is self-compressing type, further tightening of the screws will displace the bone ends together until they become compressed. After completion, the wound is irritated with sterile saline to remove tissue debris and then closed in two layers. More stable fixation is achieved when the bone ends are permitted to come into contact with each other. This reduces the bending moment experienced by the plate. Longer plates with multiple holes should be considered as well for heavier or more active animals.       

Figure 4.  Dynamic bone plates applied to the metatarsal fracture (13)

 

8. Applications

 

8.1. The mandibular reconstruction plate

There have been many researches on reconstruction plates for mandibular fractures.

Rapid rehabilitation is the expected benefit of using a bridging bone plate after composite resection for oral malignancy involving the lateral mandible (Figure 5). A bridging plate covered by a healthy myocutaneous flap is a reliable and effective method of primary reconstruction in high-risk patients with advanced cancer and uncertain long-term survival. Plates permit restoration of speech, mastication, swallowing, and facial contour. Titanium plates do not interfere with planned radiotherapy. Secondary bone reconstruction is made easier with a well adapted bone plate which provides fixation and a durable contour (14).

Figure 5.  Drawing of custom-made reconstruction plate with acrylic spacer attached medially by titanium screws (14)

When a segmental defect of the mandible is reconstructed primarily by a vascularised bone graft, there is an increased complication rate and prolonged hospitalisation. This is probably related to a longer operative period with greater blood loss and the need for delayed feeding by percutaneous gastrotomy or nasogastric tube. Younger patients with good general status and a reasonable probability of survival may benefit from primary vaseulariscd bone reconstruction, thus avoiding a second operation. Some very experienced surgeons still reserve both vascularised and free bone grafts for secondary reconstruction in cancer patients who have withstood the test of time.

In another study concerning on mandibular fractures , the goal was to retrospectively evaluate the use of 2.4-mm AO titanium reconstruction plates (15).

Patients and Methods: 

The clinical and radiologic data of 63 patients with 63 single fractures (53 comminuted, 5 dislocated, and 5 with bone loss) and 2 patients with double fractures were analyzed . Fracture location was symphysis in 37 patients (56.9%), body in 13 (20%), and angle in 15 (23.1%). The mechanism of injury, time between injury and surgery, gender and age, temporary maxillomandibular fixation (MMF) and its duration, and surgical approach was recorded. Follow-up examinations were performed at 1, 3, 6, and 12 months, at which time they noted the status of healing and any complications.

Results:

Fifty patients (77%) had a successful treatment outcome without complications; 13 patients (20%) developed minor complications; and 2 patients (3%) developed nonunion with infection requiring hardware removal and reosteosynthesis with bone graft.

Conclusions:

 It was found that 2.4-mm AO titanium reconstruction plates can be used to treat severe mandibular fractures with a low rate of major complications (3%) and a high success rate. The Figures 6,7 and 8 illustrates the application of plates and screws to the fracture site.

Figure 6.  The two plates and screws placed along the fracture site (16)

Figure 7. The pre-operative radiograph of a patient with a mandibular fracture (16)

Figure 8.  The three month postoperative radiograph after healing and plate removal (16)

8.2. The development of biodegradable screw-plate systems for maxillofacial surgery:

During the past decades the operative treatment of maxillary fractures has been influenced and modified by a variety of experimental and clinical results. The most common used systems are made of titanium and stainless steel. Disadvantages of metallic fixation are the second operation to remove the implant, the atrphic changes of the underlying bone and the possibility of allogenic response. 

The possibility of maxillary fracture treatment by use of biodegradable plates and screws made of Poly (L-lactide) has been previously shown in animal and clinical studies by several authors.These studies showed that Poly (L-lactide) has sufficient initial strength as an osteosyntheses material for maxillofacial surgery but lacks of acceptable degradation kinetics (6).

The degradation rate of Poly (L-lactide) depending on the molecular weight is between 2 to 4 years. The biocompatibility of biodegradable polymers is related to their mass loss over the time. Therefore it is very important to develop and design biodegradable implants with as small as possible dimensions to give no foreign body reactions. This is in contrast to their mechanical properties because they are much lower compared to the one of metal implants. 3 different screw plate systems were designed and tested and two different polymers were used. To improve the force transmission of the systems the screw head was welded to the plate

In the study of Timothy A. Turvey et.al.(17) the authors' experience with self-reinforced biodegradable bone plates and screws to stabilize maxillary and mandibular osteotomies was described. Patient acceptance, demographics, types of osteotomy, means of stabilization, etiology of the deformity, complications, and patient disposition are reviewed.

Patients and Methods:

Seventy patients underwent 194 osteotomies of the maxilla and/or mandible. Stabilization of each osteotomy was achieved using self-reinforced polylactite bone plates and/or screws of similar size and configuration to that of titanium systems. Placement of the devices was accomplished transorally and transfacially, consistent with the osteotomy approach. Maxillomandibular elastics were used to control the position of the jaws in each patient.

Results:

There was good patient acceptance of the material (70/74). Stabilization was accomplished as planned in all patients. Three patients experienced problems that resulted in immediate loosening of the bone screws. The remaining 67 experienced no short-term problems (6 to 24 months), and healing progressed uneventfully. In each case, acceptable occlusion and favorable aesthetic changes were noted.

Conclusions:

The experience with self-reinforced polylactite bone plates and screws to stabilize maxillary and mandibular osteotomies has been favorable on short-term observation.

 

8.3. Fractures of the frontal bone

Fractures of the frontal bone and frontal sinus are relatively uncommon. Various methods for their management have been suggested, depending on the site and extent of the damage and whether the fracture involves the anterior wall of the frontal sinus alone or is combined with an injury to the posterior wall with its consequent increase in the risk of damage to the meninges, or frontal lobes, or both.1,2 Computed tomography (CT) has simplified the management of these injuries considerably by allowing a more confident diagnosis of the extent of the injury to be made.

Where fractures of the posterior wall and tears of the dura are present, they require dural repair with or without cranialization of the sinus, or obliteration of the sinus using various autologous or alloplastic materials.3 Depressed fractures of the anterior wall are usually treated by elevation and internal fixation of the bone fragments either with stainless steel wire

or titanium/vitallium microplates

8.4. A biomechanical study of a non-contact plate system

Based on existing knowledge of noncontact plates, an experimental prototype of a nonperiosteal contact internal fixation implant ("noncontact internal fixator") has been designed by Karnezis et.al.. The construct rigidity of osteotomised synthetic composite femora, fixed with the noncontact fixator and a reamed, statically-locked intramedullary nail were compared in axial compression, two-plane bending and torsion in four types of diaphyseal fractures. With the exception of axial loading in the presence of extensive comminution, the fixation stability provided by the noncontact fixator is significantly higher than that of the tested intramedullary nail. Any degree of cortical contact between the two main fragments is important for the stability of this nonperiosteal contact fixation system under axial load. Appropriately-designed "internal fixators" could provide not only a number of biological and technical advantages, but also fixation stability comparable and in certain aspects superior to that of other fixation methods (18).

Internal fixation, based on anatomical reduction and rigid fixation of diaphyseal bone fragments, is no longer the goal in surgical fracture treatment. Minimal surgical trauma and preservation of the blood supply of bone and tissues are considered the most important factors if complications such as delayed union, non-union, infection or refracture are to be minimized. Both low-contact and non-contact plates have been used for what is referred to as “biological osteosynthesis”; long conventional plates are used for subfascial non-contact fixation in order to minimize damage to viable tissue.

Based on previously-published mechanical woks, in this research a protype of a non-contact internal fixator has been designed. Important features in the design were the wide spread of 6 mm screws, the use of low-profile screw heads and the use of a self-locking, self-preloading screw systems.The aim of this study was to investigate the extent to which a nonperiosteal-contact fixation system could provide satisfactory fixation stability for clinical use in load-bearing bones. To this end, the initial mechanical stability of the experimental non-contact internal fixator prototype and an AO universal locking nail have been compared.

8.4.1.Material and Methods

The noncontact internal fixator plate was 320 mm long, 5 mm thick and 20 mm wide as seen in Figure . There were 6 threaded holes spaced at 50 mm apart with a 80 mm hole-free zone in the middle off the plate. Only 4 holes were used for screws in the tests.

Figure 9. Non-contact internal fixator

8.4.2. Specimen Preparation

Glass-reinforced plastic composite femora were used to eliminate variations in geometry and material properties. This femur model offers the physical properties of real bone and was used to reduce the effect of interspecimen variability of cadaveric material.

8.4.3.Mechanical Testing

The models were tested in axial compression using an Instron 4302 universal testing machine. Axial torsion and cantilever bending were also tested using special rigs. The maximum applied loads were 400 N(axial compression), 16 Nm(cantilever bending), and 10 Nm (axial torsion).

8.4.4.Results:

Structural rigidity group comparisons for the tested loading conditions in terms of increasing severity of fracture pattern are given in Figure 10.

Figure 10.  Fracture Types

The mean medio-lateral rigidity of the intact composite femur was 335 Nm². The non-contact internal fixator had a mean medio-lateral flexural rigidity of 533 Nm²  (type 1 fracture) gradually decreasing to 189 Nm²(type 4 fracture).The mean medio-lateral rigidity of the non-contact internal fixator was higher than that of of the universal femoral nail for all fracture types tested.

The mean torsional rigidity of the intact composite femur was 234 Nm². The torsional rigidity of the intact femur fixed with non-contact internal fixator was higher than that of the universal femoral nail.

From these data given above, appropriately designed internal fixators may provide a number of biological and technical advantage as well as fixation stability. Further research into the biological and mechanical parameters of this concept is needed. 

8.5. Performance study of braided carbon/PEEK composite compression bone plates

In addition to unidirectional laminates and short fiber reinforcements for compression bone plate developments in the literature, using a textile structure, i.e. braid preform, have been proposed for this purpose. In the present paper, the influence of braiding angles and plate thicknesses on the bending performance of the braided composite bone plates is investigated. As a result, the influence of the braiding angle, varied in a certain range, on the plate bending properties is not significant when the plate thickness is thin. This influence becomes higher with an increase in the plate thickness. A 10° braiding angle has been seen to be appropriate for all the cases under consideration. The present study indicates that the braided composite plate with 2.6 mm thickness can be suitable for forearm treatment whereas the braided composite plate of 3.2 mm thickness is applicable to femur or tibia fixation (19).

Patients with diaphyseal fracture of a long bone are treated by the internal fixation device, such as a compression bone plate. The conventional compression plates are made of metal materials that have 5–10 times higher modulus than cortical bone. Using metal plates, it normally takes 1–2 years to obtain a complete bone healing. The mismatch between the moduli of the metal plates and the cortical bone leads to a situation where the majority of the load is transferred by the plate rather than by the underlying bone. In order to avoid "stress-shielding effect", it is desirable to use plates whose mechanical properties are close to those of cortical bone. In addition to their over high modulus, metal materials may show relatively weak strength under cyclic loading. Compression bone plates are normally subjected to extremely high cyclic loads. These examples suggest that the materials proposed for compression bone plates must also possess sufficiently high fatigue strength. Another shortcoming of metal plates is that the healing status of the fractured bone fixed beneath the plate may not be correctly identified due to metal's radio opaque, which in some cases results in undesirable artifacts in X-ray radiography.

In order to resolve these problems, polymer based composite materials, have been proposed for bone plate fixations as alternative of metal materials. However, the composite plates developed in the past mainly adapted UD (unidirectional) laminates and discontinuous short fibers as reinforcement. Environment of a compression bone plate is greatly severe because of high fatigue loading in body fluids. Although composite plates made of discontinuous short fibers have the advantage of easy fabrication by injection molding, the plates need large thickness to avoid fracture. On the other hand, compression plates made of UD laminates have to be drilled to make screw holes, resulting in a reduction of their load carrying capacity due to the breaking down of the fiber continuity. With regard to these drawbacks, the present authors have suggested using textile reinforcement to make composite compression bone plates in recent years.

 The advantage of the textile reinforcement is that the plates can sustain multi-directional loads as a drilling process to make screw holes is unnecessary and hence the yarn continuity can be retained. In an initial effort, the effectiveness of braided fabric reinforcement was recognized by using a carbon/epoxy material system. However, since epoxy resins have a possibility to give harmful influence to human body, the matrix of braided composite compression plates was later replaced with a thermoplastic PEEK (poly ether–ether-ketone) material, which is known to be well biocompatible .In this work, the braided carbon/PEEK bone plates of three different plate thicknesses each with three different braiding angles were comparatively studied in terms of their bending performance.

8.5.1.Materials and Fabrication

In this study, the micro-braided yarn, which contained PEEK matrix and reinforcement carbon fibers, was used to fabricate braided carbon/PEEK compression bone plates. The unique feature of this yarn is that the reinforcing and matrix fibers are easily mixed using a simple braiding technique. Flat braided fabrics were preformed using micro-braided yarns, as indicated in Figure 11.  In terms of a hot press machine, a composite bone plate was obtained by placing multiple layers of the flat braided fabrics in a stainless-steel mould to which pins were attached to form screw holes without breaking the yarn continuity as seen in Figure 12. Braided fabrics with three different braiding angles, i.e., 5°, 10° and 15°, were used to investigate the influence of the braiding angles on the bending performance of the composite plates. In order to further investigate such influence, three different plate thicknesses, 2.6, 3.2 and 3.8 mm, were achieved using different fabric layers.

Figure 11. Photograph of a flat braided fabric of micro-yarns

Figure 12. Schematic Drawing of insertion way of flat braided fabric

Bending behavior of a bone plate is one of the most critical mechanical properties from an application viewpoint, and is generally evaluated by maximum bending moment and bending stiffness calculated from the initial linear moment against the total bending angle (angulation) of the plate. In this study, static four-point bending tests were conducted with a cross-head speed of 1.0 mm/min at room temperature

8.5.2. Results

Braided composite compression plates

Angulation curves of the braided composite compression plates with three different thicknesses were calculated. In the case of the specimens with 2.6 mm thickness, in all of   bending moment increased linearly with angulation regardless of different braiding angles. When the bending moment reached around 6 N m, some plastic deformation began to occur and all the specimens finally fractured at the third hole near a loading point. On the other hand, all the other specimens with larger thicknesses (i.e., 3.2 and 3.8 mm) did not show a significant plastic behavior.  In other words, the bending moment increased essentially linearly until rupture and all the specimens finally fractured in the same way as those with the minimum thickness.

Performance comparison

The braided composite plates showed an increase in the maximum bending moment with the increase of plate thickness for all the braiding angles. The plates which have the minimum thickness showed only 8% difference in the bending moment among different braiding angle specimens. This difference was more distinct for the thicker specimens, i.e., 18% for the 3.2 mm and 19% for the 3.8 mm thick specimens among different braiding angles. The yield bending moment of the stainless-steel plate was relatively lower than its maximum bending moment. However, the braided composite plates did not show drastic yield moment decrease as compared with their maximum value. It is noted that the braided composite plates of 3.2 mm thickness showed almost the same yield moment as that of the stainless-steel plate which was 3.8 mm thick. The bending stiffness of the braided composite plates increased with the increase in thickness, and the specimens of 3.8 mm thickness indicated a quite close stiffness to that of the stainless-steel plate. There was no stiffness variation for the 2.6 mm thick specimens with the three different braiding angles.

8.5.3. Conclusion

Although several composite bone plates were developed using UD laminates and discontinuous short fibers to serve as alternatives for the conventional stainless-steel AO compression plates, there still remain a number of improvements to be addressed from a mechanical viewpoint. In this regard, using braided carbon/PEEK fabric composites have been proposed as a new material system for bone plate development. In this work, the influence of braiding angle and plate thickness on the bending properties of the braided composite bone plates is further investigated. As a result, the influence of the braiding angle, varied in a certain range, on the plate bending properties is not significant when the plate thickness is thin. This influence is higher when the plate becomes thicker. A 10° braiding angle can be recommended for all the cases under consideration. It is considered that the braided composite plate with 2.6 mm thickness can be suitable for forearm treatment whereas the braided composite plate of 3.2 mm thickness is applicable to femur or tibia fixation.

9.Some illustrations of bone plate applications:

Figure13: Fracture of the middle phalanx repaired by bone plate application to the front of the pastern.  This enables lag screw fixation of fragments to the plate and compression to fuse the middle phalangeal joint (16)

Figure 14. Bone plate stabilization and microvascular anastomosis of free fibula graft reconstruction of the mandible (16)

Figure 15.  Repair of an ulnar fracture with a bone plate used as a tension band on the caudal aspect of the ulna (16)

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