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Compare osseoincorporation and osteosynthesis.

Osseointegration (from Latin ossum "bone" and integrare "to make whole"), sometimes also referred to as osteointegration, is the direct structural and functional connection between living bone and the surface of a load-bearing artificial implant ("load-bearing" as defined by Albrektsson et al. in 1981). Osseointegration played an important role in enhancing the science within the medicalbone and joint replacement techniques as well as dental implants and improving prosthetics for amputees.


Osseointegration is defined as "the formation of a direct interface between an implant and bone, without intervening soft tissue".[1]

Osseointegration may also be defined as:

  1. Osseous integration, the apparent direct attachment or connection of osseous tissue to an inert alloplastic material without intervening connective tissue.
  2. The process and resultant apparent direct connection of the endogenous material surface and the host bone tissues without intervening connective tissue.
  3. The interface between alloplastic material and bone.

A more recent definition (by Schroeder et al.) defines osseointegration as "functional ankylosis (bone adherence)", where new bone is laid down directly on the implant surface and the implant exhibits mechanical stability (i.e. resistance to destabilization by mechanical agitation or shear forces).

History[edit]

Osseointegration was first observed—albeit not explicitly stated—by Bothe, Beaton, and Davenport in 1940.[2][3] Bothe et al. were the first researchers to implant titanium in an animal and remarked how it tended to fuse with bone.[2][3] Bothe et al. reported that due to the elemental nature of the titanium, its strength, and its hardness, it had great potential to be used as future prosthesis material.[2][3] Osseointegration was later described by Gottlieb Leventhal in 1951.[2][4] Leventhal placed titanium screws in rat femurs and remarked how "At the end of 6 weeks, the screws were slightly tighter than when originally put in; at 12 weeks, the screws were more difficult to remove; and at the end of 16 weeks, the screws were so tight that in one specimen the femur was fractured when an attempt was made to remove the screw. Microscopic examinations of the bone structure revealed no reaction to the implants. The trabeculation appeared to be perfectly normal."[2][4] The reactions described by Leventhal and Bothe et al. would later be coined into the term "osseointegration" by Per-Ingvar Brånemark of Sweden. In 1952, Brånemark conducted an experiment in which he utilized a titanium implant chamber to study blood flow in rabbit bone. After the experiment, when it was time to remove the titanium chambers from the bone, he discovered that the bone had integrated so completely with the implant that the chamber could not be removed. Brånemark called this "osseointegration", and, like Bothe et al. and Leventhal before him, saw the possibilities for human use.[2][3][4]

Headshot of Per-Ingvar Brånemark

In dental medicine the implementation of osseointegration started in the mid-1960s as a result of Brånemark's work.[5][6][7][8] In 1965 Brånemark, who was at the time Professor of Anatomy at the University of Gothenburg, placed dental implants into the first human patient – Gösta Larsson. This patient had a cleft palate defect and required implants to support a palatal obturator. Gösta Larsson died in 2005, with the original implants still in place after 40 years of function.[9]

In the mid-1970s Brånemark entered into a commercial partnership with the Swedish defense company Bofors to manufacture both dental implants and the instrumentation required for their placement. Eventually an offshoot of Bofors, Nobel Pharma, was created to concentrate on this product line. Nobel Pharma subsequently became Nobel Biocare.[9]


Brånemark spent almost 30 years fighting the scientific community for acceptance of osseointegration as a viable treatment. In Sweden he was often openly ridiculed at scientific conferences. His university stopped funding for his research, forcing him to open a private clinic to continue treating patients. Eventually an emerging generation of young academics started to notice the work being performed in Sweden. Toronto's George Zarb, a Maltese-born Canadian prosthodontist, was instrumental in bringing the concept of osseointegration to the wider world. The 1983 Toronto Conference is generally considered to be the turning point, when finally the worldwide scientific community accepted Brånemark's work.


In the past decades osseointegration has emerged further and today osseointegration is a predictable and commonplace treatment method.[9] It has become an area of both lab research, clinical trials, and has received FDA approval for clinical usage, where some historical milestones are listed below:

  • Hydroxyapatite-coated (HA) hip implants were used for the first time in 1985.[10] More than 30 years later HA-coatings are still used to augment the cementless fixation of implants. The success of this material is attributed to that it is very similar to the mineral phase of the bone which is made of 70% hydroxyapatite and about 30% collagen. This has been proven to be biocompatible, bioactive and osteoconductive which improves the primary stability and osseointegration of implants.[11]
  • More recently studies such as those published by Al Muderis in 2010 in Sydney Australia utilized a high tensile strength titanium implant with a high prose plasma sprayed surface as an intramedullary prosthesis that is inserted into the bone residuum of amputees and then connects through an opening in the skin to a robotic limb prosthesis thereby allowing amputees to mobilize with more comfort and less energy consumption. Al Muderis also published the first studies of combining osseointegration prosthesis with joint replacement enabling below knee amputees with knee arthritis or short residual bone to mobilize without the need of a socket prosthesis.[12]
  • The first clinical studies on titanium coated PEEK-Carbon spinal cages were performed in 2012 on 42 patients. The "paper" attributed the success of the trial to the similarity in Young's modulus of the implanted material to the cortical bone at the implant site. PEEK itself is an inert polymer with a hydrophobic surface, which is not favorable for osseointegration. These drawback was overcome by coating the PEEK surface with (pourous) titanium to significantly enhance the osseointegration of the material, making it a suitable candidate for spinal implants.[13]
  • In recent years many percutaneous osseointegrated implants that are used to directly attach the prosthetic limb to the residual limb of a person with an amputation have come to the attention of FDA in the United States of America. On December 7th 2015, the first two persons in the U.S., veterans Bryant Jacobs and Ed Salau, received a percutaneous osseointegrated prosthesis. Doctors at the Salt Lake Veterans Affairs Hospital embedded a titanium stud in the femur of each patient. Two months later on February 8th 2016 they attached the docking mechanism for the prosthesis in a second operation. These veterans are part of a feasibility study which includes a 10-year trial with 10 participants. [14] A percutaneous osseointegrated implant increases life quality (compared to conventional prosthesis) significantly, however there seems to be an increased risk for infections at the implant - soft tissue interface. Most of the time thees infections can be treated successfully with antibiotics.[15] These types of implants are considered Class III devices and require the highest degree of control to assure that they are safe and effective. However they have been approved by the FDA through the Humanitarian Device Exemption pathway and can be used in a small number of patients per year.[16]

Mechanism[edit]

See also: Titanium biocompatibility

Osseointegration is a dynamic process in which characteristics of the implant (i.e. macrogeometry, surface properties, etc.) play a role in modulating molecular and cellular behavior. Both calcium phosphate coated implants and titanium implants are stabilized chemically with bone, either through direct chemical bonds between calcium and titanium, or by the bonding to a cement line-like layer at the implant/bone interface.[17][18]

Stages of Osseointegration[edit]

Titanium implant (black) integrated into bone (red), with fibrous connective tissue (yellow)

The mechanism of osseointegration is similar to direct bone healing and primary fracture healing though traumatization around the bone, which activates the osseointegration process.[19][20] The osseointegration process includes three biological stage.[21][22]

  • Incorporation by woven bone formation
  • Adaptation of bone mass to mechanical load
  • Adaptation of bone structure to mechanical load.

Incorporation by Woven Bone Formation[edit]

Bone healing caused by fractures or by incorporation of implants is activated when the pre-existing bone matrix suffers from lesions, where the primary formation of bone can be divided into three main phases: inflammatory phase, proliferative phase and maturation phase.[23]

The inflammatory phase is similar to general wound healing, and starts when an implant is placed in contact with bone. The contact (and lack of perfect contact) between the implant and bone tissue causes reactions in the osseous defect area such as protein adsorption, blood coagulation, thrombosis and cytokinesis. It also triggers the cellular generalized inflammatory response, causing migration of white blood cells to the affected area.[23]

The proliferative phase starts within the inflammatory phase, and is characterized by the creation of blood vessels. Vascular ingrowth formation (neovascularization), as well as formation of an oxygen gradient occurs by the wound edge as a result of the metabolism of local inflammatory cells. The oxygen deficit close to the lesion, combined with certain cytokines, stimulates further vascular growth (angiogenesis). The cells are also responsible for the development of the extracellular matrix at this stage, as well as initial fibrous tissue, allowing fibrocartilaginous callus formation as a base substance for final bone formation which then moves on into the maturation phase.[23]

In the maturation phase, where the actual bone formation starts, new bone is constructed to replace the necrotic bone created from the inevitable implant damage.[23] Woven bone grows on the scaffold of dead bone trabeculae, and at the same time the fibrocartilaginous callus forms, eventually resulting in a primitive type of bone tissue. The woven bone formation is characterized by a random, felt-like orientation of collagen fibrils, osteocytes, and a relatively low mineral density. This primary formed bone however has a great capacity for growth. It generates a vascular net at a rapid rate, forming spongiosa that can bridge gaps of almost 1 mm in a couple of days. This sponge-like structure lends itself well to use it as a basic structure for the further formation of the bone. The bone mainly starts to grow from the surrounding bone towards the implant, except for in narrow gaps where it concurrently forms upon the implant surface.[24]

Adaptation of Bone Mass to Load[edit]

After the first month of implant integration, the microscopic structure of the newly formed bone changes towards the parallel-fibered bone and the lamellar bone. Appositional growth of lamellar bone in parallel layers with fibers in alternating directions creates a strong structure (similar to plywood), where the parallel-fibered bone acts as an intermediary layer between the woven bone and the lamellar bone. Neither type of bone can form the woven bone scaffold structure. The growth is solely determined by apposition on a preformed solid base, thus limiting the growth to: the woven bone formed, pre-existing bone surface or the implant surface.[24]

  • Mature bone deposition on woven bone results in a concentrated bone reinforcement of areas where major forces are transferred from the implant to the original, surrounding bone.
  • Deposition of bone on a pre-existing bone surface compensates for the necrotic trabecular bone caused by trauma and interruption of blood supply at surgery, reinforcing the bone loss, and may reflect the preferential strain pattern resulting from functional loads.
  • Lastly, the bone deposition on an implant surface increases at the bone-implant intercept, and thus distributes the external load over a larger area. This is considered to represent the adaptation of bone-mass to load due to the growth response on the load magnitude and direction.[24]

Adaptation of Bone Structure to Load[edit]

The last stage of osseointegration is the adaptation of bone structure to load, which includes bone modeling and remodeling. It starts around the third month after an implant insertion, firstly with a high level of activity that later slows down, but still continues for the implant´s life. The remodeling occurs in discrete units for both cortical and cancellous bone, which are also referred to as bone multicellular units.[24] The remodeling occurs through osteoclast resorption followed by lamellar bone deposition, where there needs to be both resorption and deposition of bone for a stable growth. Functional adaptation of bone is based on shape deformation and a coupled resorption and formation rate.[24] Generally, remodeling in osseointegration and adaptation of bone structure to load happens in two ways:

  • It improves bone quality by replacing pre-existing necrotic bone and/or initially formed, more primitive woven bone with mature, viable lamellar bone.
  • It leads to a functional adaptation of bone structure to load by changing the dimension and orientation of the supporting elements.
Modification of Bone Growth Integration with Drugs[edit]

The adaptation of bone structure to load can be modified by drugs that alter the bone integration process. Bisphosphonates drugs commonly used in clinical applications in this area.[25][26][27] One bisphosphonate drug is zoledronic acid (ZOL), which improves the osseointegration by blocking osteoclasts, thus inhibiting the resorption of old bone, and allowing the growth rate to be more significant than usual.

Implant Surface Characteristics[edit]

Protein Adsorption & Coagulation[edit]

Each implant comes into contact with blood and tissue immediately after implantation and is covered by a layer of proteins and glycoproteins. At the beginning, the proteins bind loosely to the surface. This so-called "conditioning film" provides an attack surface for the adhesion of other proteins and cells. This film has clumping factors that bind fibrinogen and fibrin, which in turn promote cellular adhesion to the biomaterial.[28] However, medical implants bear the risk of triggering infections and the formation of fibrosis capsules when they come into contact with blood. The formation of too thick fibrosis capsules, can lead to, for example, the failure of implants. The problem is that specific tissue reactions to different surface properties of implants have not yet been fully clarified. Even an inflammatory response to inert, non-immunological and non-toxic materials is difficult to avoid. It is known so far that after implantation, proteins accumulate around the implant and lead to the accumulation of inflammatory cells. Therefore, adsorbed proteins play an important role in the immune response and coagulation cascade. The goal of current research is to control the amount and composition of endogenous proteins that attach to the surface, and the degree of conformational changes in proteins. The problem is that conformational changes after attachment to the surface epitopes can present binding sites for inflammatory mediators and thus start inflammation. The interaction between proteins and surfaces depends on both the protein properties and the surface properties. The goal can to be achieved through different surface modifications, which have been the subject of intensive research over the last 20 years.[29] Temperature, pH value, ionic strength, and buffer composition are all properties which can influence the protein adsorption rate too.[30] Nowadays, biomaterials are mostly hydrophobic and have a high affinity to a wide range of proteins, as they are mostly hydrophobic themselves. Proteins such as plasma proteins, albumin, fibrinogen, IgG, fibronectin and Von Willebrand factor often accumulate after implant placement. Hydrophobic interactions can lead to the presentation of hydrophobic domains of proteins, which allows them to attach to the surface. The exception in this case are the less hydrophobic glycoproteins that are preferentially adsorbed on hydrophilic surfaces.

These conformational changes could lead to the initiation of different reactions such as inflammation, coagulation or foreign body reaction. The assumption is that the biomaterials interaction and the subsequent conformational change of the proteins uncovers hidden protein structures and/or sequences that interact with inflammatory cells and can induce or enhance the resulting foreign body reaction. A recognized reason for the conformational changes of proteins is the difference of free energy in solution and on surfaces. This difference and may therefore lead to a conformational change on the surface due to lower free energy.[31] As previously mentioned, surface properties can affect protein adsorption, denaturation, and epitope presentation. Typical characteristics for "shedding" proteins from the surface are: hydrophilicity, the presence of H-bridge acceptors, non-H-bond donors, and a neutral charge. The identification of functional groups that repel proteins can help to "tune" material surfaces, control protein adhesion and its interaction, and increase the associated biocompatibility.

Influence of surface functional groups cellular responses[edit]

The most common functionalities investigated in relation to biomaterial interactions are the carboxyl (-COOH), hydroxyl (-OH), amino (-NH2), and methyl (-CH3) groups.

  • Carboxyl functional group-bearing surfaces: Studies have shown that biomaterial surfaces carrying carboxyl groups have a negative charge that causes fibronectin and albumin to elute better. Surfaces with a carboxyl group also show increased cell growth. However, the properties depend on the concentration of this functional group, since too high a concentration results in an increased density of this functional group and a low cell growth.
  • Hydroxyl functional group-coated surfaces: Hydroxyl groups on the surface tend to have a neutral charge. Adsorbed fibronectin on the hydroxyl groups shows increased cell adhesion strength and, compared to other functional groups, show hydroxyl functional groups in experiments with osteoblasts a higher level of differentiation and mineralization. The disadvantage is that it has low compatibility with plasma proteins and a low associated platetet compatibility.
  • Amino functional group-rich surfaces: Compared to the carboxyl group, surface-coated materials with amino groups have a positive charge. Fibronectin and osteopontin show favorable protein conformations after adsorption, which positively influences endothelial cell growth through higher presentation of binding receptors and focal adhesion components.
  • Methyl functional group-bearing surfaces: The methyl group makes up the main constituent of polymer materials has a hydrophobic surface. The hydrophobic surface caused by the methyl group increasingly adsorbs proteins but this typically causes an adverse conformational change that is unfavorables to new cell interactions. One result shows increased fibrinogen binding, platelet accumulation and low blood compatibility. Another study shows that fibrinogen, albumin and IgG bind more heavily to such a surface. Nevertheless, the results show undesirable surface reactions with cells.
  • Mixed Functionality surfaces: The goal with these types of surfaces is to use a combination of different functional groups to refine the favourable properties of both functionalities and to improve biocompatibility.[32]

Because protein-surface interaction experiments can take a long time, be costly and are mostly tested in-vitro, computer simulations are being developed today to better understand protein-surface interactions. This allows the movement or behavior of each protein and atom to be accurately tracked. As this area of ​research is very complex and extensive, and there are often new problems, new methods need to be developed to better understand the functionality of surfaces and protein-surface interaction.

The effort on this topic is still rewarding as it clearly influences the biocompatibility of the implants. Thus, the healing after implantation can be significantly improved and shortened. This has advantages for both the patient and the hospital as the likelihood of infection is reduced and the hospitalization period is shorter.

Nanotopography[edit]

Nanotopography plays an important role in the process of osseointegration. A good nanotopography can promote bone bonding and strengthen the interface, which is essential for the osseointegration process.[33]

Methods to obtain a nanostructured implant surfaces include:

A suitable nanostructured surface has shown to increase the alkaline phosphatase synthesis and increase calcium mineral content in the cell layer, and it can also reduce the risk of developing a fracture, furthermore it offers a greater bone contact area.[34]

The main reason why the nanotopography affects the osseointegration of implant in such a way is because structures that are smaller than 100 nm resemble more closely the constituents of the bone. Also an implant that has a nanostructured surface has a far greater surface area then an implant with a microstructured surface, allowing the bone cells to better attach to the surface of the implant and to interact with it.[34]

Nanostructures present on the surface of an implant can also reduce the inflammation response after a surgical procedure. In rat models it has been found that the number of macrophages present at the surface of the implant was smaller in the case of the nanostructured implant compared to a traditional implant. The decrease in the number of macrophages is attributed to a decrease in the secretion of the TNF-α, which is the major proinflammatory and chemotactic cytokine. However this effect is closely related to the shape and size of the nanostructures. Nano semi-spheres and nano tubes decreased the production of TNF-α, MCP-1, MIP-1α, IL-1β and IL-6, while nano grooves stimulate the secretion of IL-1β and TNF-α. Also the size of the nanostructures has an effect on the secretion of inflammatory factors. It was observed that 30 nm nano tubes induce less inflammatory cell infiltration then the 80 nm ones.[35]

Mechanical Characteristics of Natural Bone Imitating Implants[edit]

Often long term implants in bones are designed to take over the function of the lost or broken bone. It is therefore a reasonable idea to replace the bone with a material with similar mechanical properties to healthy bone. The human cancellous bone possesses a stiffness ranging from 12-23 GPa.[36] However, the stiffness values of bulk metal and technical ceramics are one order of magnitude higher than that of human cancellous bone. By creating foams out of bulk materials, similar stiffness values to cancellous bones can be reached. However predicting the stiffness of foams is not trivial. A suitable stiffness of the implants is important for a good osseointegration because a mismatch of stiffness between the implant and the bone results in stress-shielding effects.[37] Then the implant, which is typically stiffer than the bone, will absorb most of the load. According to Wolff's law only bone that feels a certain minimal load will grow and bone without load will slowly disappear. Thus a stiffness mismatch can lead to bone density reduction and a loosening of the implant, leading to an implant failure.[38]

Porous Osseointegrating Implants - Ideal Characteristics[edit]

The ideal implant for osseointegration should be porous biocompatible material and have similar mechanical properties to the host bone in the surrounding area while providing sufficient mechanical strength (Wolff's law). The porous implant should have interconnected pores with a size from 200 um to 600 um. Such a porosity implant will allow sufficiently easy fluid flow for cell nutrition and osteoblast multiplication as well as migration for cellular colonization of the implant to become uniform.[39] To increase the surface area, which also facilitate vascularization and the ingrowth of new bone[40], the porous scaffold should have a rough nano-porous coating.The physical strength of the implant should be high enough to survive surgery as well as the in vivo forces. This can be a limiting factor for the porosity of the implant, as porosity reduces the physical strength of the implant.

To achieve repeatable results, a repeatable production of the implant is needed. This got only recently possible, due to the advanced in additive manufacturing which allows a create complex reproducible 3D-structures with bulk porosity not only a porous coating. At the same time the ideal implant should be reasonable priced and reproducible implantable. So far no such product exists.[41]

Implant Material Classes and Their Affect[edit]

Metals[edit]

Metals are the most used materials in implant devices, due to their properties, such as the elastic modulus, hardness, fracture toughness and wear resistance; which are critical to the performance of an implant. The most employed metallic biomaterials are titanium (Ti) and titanium alloys, cobalt chromium (CoCr) alloys, and stainless steel. The surface composition and structure of metal implants is important for osseointegration mechanism.[42] Titanium alloys show a more rapid integration, compared to other alloys. Oxidation of Ti surfaces and formation of a surface TiO2 layer through the process of passivation, generate a significantly stronger interface, that is essential to improve osseointegration. Indeed, it has been demonstrated that a thicker TiO2 layer improves surface wettability and cell adhesion. Furthermore, apatite crystals can grow on the surface of TiO2, induced by OH- groups in the oxide layer.[43]

Chemical and biochemical modification of metal surface:[44] Ti and CoCr alloys are biocompatible, but not bioactive. By manipulating the chemical and biochemical composition of the material, bioactivity can be induced. There are two main method: coating and impregnation. Surface coating with inorganic molecules, e.g. calcium phosphate, increases osteointegration. Plasma spray coating with hydroxyapatite (Ca10(PO4)6(OH)2), remains one of the most common methods of surface modification of clinical implants due to the similarity of hydroxyapatite to the mineral phase of bones. Surface impregnation with Mg, S, P and Ca has been reported to drive osseointegration and improve bone to metal contact in vivo. Another important method is the bio-functionalization of the surfaces with organic molecules. These modifications include linking of peptides and proteins, ECM proteins, growth factors, and pharmacologically active molecules. These surface functionalization increase protein adsorption and cell adhesion, improving the osseointegration.

Metal Foams[edit]

See also: Titanium foam

Since 2005, a number of orthopedic device manufacturers have introduced products that feature porous metal construction.[45][46][47] Clinical studies on mammals have shown that porous metals, such as titanium foam, may allow the formation of vascular systems within the porous area.[48] For orthopedic uses, metals such as tantalum or titanium are often used, as these metals exhibit high tensile strength and corrosion resistance with excellent biocompatibility. By making a porous structure, the stiffness (young's modulus) of the bulk metal can be reduced and adapted to the stiffness of the natural bone leading to no/lower stress concentrations.

SEM picture of a metal foam

The process of osseointegration in metal foams is similar to that in bone grafts. The porous bone-like properties of the metal foam contribute to extensive bone infiltration, allowing osteoblast activity to take place. In addition, the porous structure allows for soft tissue adherence and vascularization within the implant. A study about titanium foam showed that the bone exhibits its natural ability to adjust local fiber away from the low-stress regions toward high stress regions.[49] Such materials are currently deployed in hip replacement, knee replacement and dental implant surgeries.

Recent advancements in additive manufacturing (e.g. electoron beam and selective laser sinthering) allow the repeatable manufacturing of exact geometries with a known size and distribution of pores, leading to an interconnecting sponge structure with sub-micron sized pores.[39]

Ceramics[edit]

Ceramics are very common materials used in bone implants. Bioactive ceramics can form bone bonds through a carbonate apatite layer.[50] Especially, the use of Hydroxyapatite (HA) ceramics is relevant, due to the similarity with the mineral phase of bone matrix. The grade of porosity and the pores size of the material are important factors, which could increase the protein adsorption and cell adhesion, and therefore also the osseointegration.[51] Nevertheless, the strength of ceramics decreases when porosity grade increases, and the brittleness of these materials restrict the orthopedic application. Despite this, ceramic materials like hydroxyapatite and bio-glass are important bioactive materials that promote bone integration, throughout the formation of calcium phosphate layer between the bone and the implant. They can be used also as a coating to increase the bioactivity of other material, such as metals.[52] Another important class of ceramic materials consist of bioresorbable calcium phosphate materials; they degrade with time in presence of physiological fluids, by active or passive resorption. Important applications are healing of bone defects, fracture treatment, total join replacement, orthopedics and bone fillings.[53]

Ceramic-Coatings[edit]

Ceramic surface coating can be deployed in order to increase bioactive properties of materials and improve osseointegration. For example, Titanium is a biocompatible material and its propensity to form an oxide layer on its surface prevent corrosion problem.[54] However Titanium has poor bioactivity and surface coating with ceramic can improve this property.[55] Recently, research is focused on materials with mechanical and structural properties similar to bone matrix, in order to improve the success of osseointegration and reduce immune response.[56] Hydroxyapatite (HA) is the most deployed material as coating material, due to its similarity to the inorganic bone matrix and the ability to produce hydroxyapatite layer between the implant and the bone. Plasma‐sprayed ceramic coating is the most used method to produce a surface coating. Other important coating method are sol-gel and electrophoretic deposition.[57] Titanium foams can be coated with HA through the method above-mentioned. It has been shown that HA-coated titanium exhibits increased interfacial strength in comparison to titanium foams without the coating.[58]

Polymers[edit]

Polymers are important material in biomedical application. They are composed by the repetition of monomers and their mechanical properties are correlated to their structure.[59] Polymer can be hydrolytically and enzymatically degradable, which make polymers a good material for osseointegration and osseoinduction.[60] The main advantage in the use of polymeric material as medical device is the possibility of amending the structure and the cross-link bond of polymers in order to adapt elastic modulus and other mechanical properties, to bones characteristic.[61] The porosity grade and surface composition of polymer materials are also important for protein adsorption and cell adhesion.[62] The porosity of polymers can be obtained with several methods that depend on the final function of the material. Copolymerization and polymer modification (such as PLGA) are important technique to obtain a more suitable material for osseointegration. Another important polymer in bone application is polymethylmethacrylate (PMMA), it is injectable and it can be used as bone cements in joint replacement surgery.[63] Polymer have many advantage, however synthetic polymeric material are not bioactive and in several case the degradation produce inflammation and swelling. To overcome this problem strategies have been developed to attach biomolecules that improve protein, cell adhesion and avoid immune response. [64]

Composite Materials[edit]

Nowadays carbon fibre-composites are not only used in airplanes and motorsports. Quite recently medical grade polymer implants (mainly PEEK) were reinforced with carbon fibers (e.g. for spine cages). Those implants have interesting properties from a medicinal standpoint. By changing the fiber orientation and the fiber content percentage, the anisotropic mechanical properties of bone can be mimicked and fine tuned. They can mechanically compete with metals and be X-ray transparent at the same time, allowing good visibility in CT operation and radiation treatments of cancer in the implant area, which would not be possible with a metal implant. However, there is a substantial problem with PEEK-CF implants. Bone cells do not like PEEK. Therefore an osseointegrative coating is added which allows bone integration e.g. a porous titanium coating.[13]

Technique[edit]

X-ray of right knee implant

Osseointegartion-related techniques are emerging more and more in the medical market as solution for bone deformities and various types of bone trauma, and a wide range of new solutions for applications are currently under research. There are several techniques already available on the market as stated above. This ranges from the most commonly known and used osseointegration process of dental implants, but other solutions include craniofacial prosthetics, amputee prosthetics and implants, knee and joint replacements(such as hips and fingers), hearing aid implantation, bone screws, spine cages etc. For all of these applications, the concept of osseointegrations is the same: to form an rigid interface layer between the implant and bone. But there are also cases where the osseointegration process is not sought for. One example of where the implant is in contact with bone but osseointegration is unwanted is for certain spinal fusion procedures. Some pedicle screws that goes into the spine are only temporary and will be removed when the healing is done, and should thus not create an intermediate layer between the screw and bone since this would make it harder, and with risk of damaging the healthy bone, to remove the screw.

Osseointegration & Periodontology[edit]

For osseointegrated dental implants, metallic, ceramic, and polymeric materials have been used,[65] in particular titanium.[66] To be termed osseointegration the connection between the bone and the implant need not be 100 percent, and the essence of osseointegration derives more from the stability of the fixation than the degree of contact in histologic terms. In short it represents a process whereby clinically asymptomatic rigid fixation of alloplastic materials is achieved, and maintained, in bone during functional loading.[67] Implant healing times and initial stability are a function of implant characteristics. For example, implants utilizing a screw-root form design achieve high initial mechanical stability through the action of their screws against the bone. Following placement of the implant, healing typically takes several weeks or months before the implant is fully integrated into the surrounding bone. First evidence of integration occurs after a few weeks, while more robust connection is progressively effected over the next months or years.[68] Implants that possess a screw-root form design result in bone resorption followed by interfacial bone remodeling and growth around the implant.[69]

Implants utilizing a plateau-root form design (or screw-root form implants with a wide enough gap between the pitch of the screws) undergo a different mode of peri-implant ossification. Unlike the aforementioned screw-root form implants, plateau-root form implants exhibit de novo bone formation on the implant surface.[70] The type of bone healing exhibited by plateau-root form implants is known as intramembranous-like healing.[69]

Though the osseointegrated interface becomes resistant to external shocks over time, it may be damaged by prolonged adverse stimuli and overload, which may result in implant failure.[71][72] In studies performed using "Mini dental implants," it was noted that the absence of micromotion at the bone-implant interface was necessary to enable proper osseointegration.[73] Further, it was noted that there is a critical threshold of micromotion above which a fibrous encapsulation process occurs, rather than osseointegration.[74]

Other complications may arise even in the absence of external impact. One issue is the growing of cement.[75] In normal cases, the absence of cementum on the implant surface prevents the attachment of collagen fibers. This is normally the case due to the absence of cementum progenitor cells in the area receiving the implant. However, when such cells are present, cement may form on or around the implant surface, and a functional collagen attachment may attach to it.

Testing procedures[edit]

There are a number of methods used to gauge the level of osseointegration and the subsequent stability of an implant. One widely used diagnostic procedure is percussion analysis, where a dental instrument is tapped against the implant carrier.[76] The nature of the ringing that results is used as a qualitative measure of the implant’s stability. An integrated implant will elicit a higher pitched "crystal" sound, whereas a non-integrated implant will elicit a dull, low-pitched sound.[77]

The resonance frequency analysis (RFA) is based on the same physical principal as the method above (resonance frequency). With the difference that you get a quantitative feedback with this increasingly implemented non-invasive diagnostic method.[76] A resonance frequency analyzer device prompts vibrations in a small metal rod temporarily attached to the implant. As the rod vibrates, the probe reads its resonance frequency and translates it into an implant stability quotient (ISQ), which ranges from 1–100, with 100 indicating the highest stability state. Values ranging between 57 and 82 are generally considered stable, though each case must be considered independently.[76]

Another method is a reverse torque test, in which the implant carrier is unscrewed. If it fails to unscrew under the reverse torque pressure, the implant is stable. If the implant rotates under the pressure it is deemed a failure and removed.[78] This method comes at the risk of fracturing bone that is mid-way in the process of osseointegration.[76] It is also unreliable in determining the osseointegration potential of a bone region, as tests have yielded that a rotating implant can go on to be successfully integrated.[79]

Applications[edit]

See also[edit]

Notes and references[edit]

  1. ^ Miller, Benjamin F.; Keane, Claire B. (1992). Miller-Keane Encyclopedia & Dictionary of Medicine, Nursing, and Allied Health. Philadelphia: Saunders. ISBN 0-7216-3456-7.[page needed]
  2. ^ a b c d e f Rudy, Robert; Levi, Paul A; Bonacci, Fred J; Weisgold, Arnold S; Engler-Hamm, Daniel (2008). "Intraosseous anchorage of dental prostheses: an early 20th century contribution". Compend Contin Educ Dent. 29 (4): 220–229. PMID 18524206.
  3. ^ a b c d Bothe, RT; Beaton, KE; Davenport, HA (1940). "Reaction of bone to multiple metallic implants". Surg Gynecol Obstet. 71: 598–602.
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[1]

Further reading[edit]

  • Zarb GA, Schmitt A (July 1990). "The longitudinal clinical effectiveness of osseointegrated dental implants: the Toronto Study. Part II: The prosthetic results". The Journal of Prosthetic Dentistry. 64 (1): 53–61. doi:10.1016/0022-3913(90)90153-4. PMID 2200880.
  • Apse P, Zarb GA, Schmitt A, Lewis DW (1991). "The longitudinal effectiveness of osseointegrated dental implants. The Toronto Study: peri-implant mucosal response". The International Journal of Periodontics & Restorative Dentistry. 11 (2): 94–111. PMID 1718917.
  • Chaytor DV, Zarb GA, Schmitt A, Lewis DW (1991). "The longitudinal effectiveness of osseointegrated dental implants. The Toronto Study: bone level changes". The International Journal of Periodontics & Restorative Dentistry. 11 (2): 112–25. PMID 1938184.
  • Barber AJ, Butterworth CJ, Rogers SN (January 2010). "Systematic review of primary osseointegrated dental implants in head and neck oncology". The British Journal of Oral & Maxillofacial Surgery. 49 (1): 29–36. doi:10.1016/j.bjoms.2009.12.007. PMID 20079957.
  • Hultin M, Gustafsson A, Klinge B (February 2000). "Long-term evaluation of osseointegrated dental implants in the treatment of partly edentulous patients". Journal of Clinical Periodontology. 27 (2): 128–33. doi:10.1034/j.1600-051x.2000.027002128.x. PMID 10703659.
  • Olivé, Jordi; Aparicio, Carlos (1990). "The periotest implant as a measure of osseointegrated oral implant stability". The International Journal of Oral & Maxillofacial Implants. 5 (4): 390–400.
  • Holmgren EP, Seckinger RJ, Kilgren LM, Mante F (1998). "Evaluating parameters of osseointegrated dental implants using finite element analysis--a two-dimensional comparative study examining the effects of implant diameter, implant shape, and load direction". The Journal of Oral Implantology. 24 (2): 80–8. doi:10.1563/1548-1336(1998)024<0080:EPOODI>2.3.CO;2. PMID 9835834.
  • Trabecular Metal Material: The Next Best Thing to BoneTM: http://www.trabecularmetal.zimmerdental.com/Implant/imp_home.aspx

External links[edit]


Category:Dentistry Category:Restorative dentistry Category:Implants (medicine) Category:Prosthetics Category:Oral surgery Category:Oral and maxillofacial surgery Category:Orthopedic surgical procedures

  1. ^ {{Cite journal | last= Liping Tang| first= Paul Thevenot| year= 2011| title= Surface Chemistry influence implant biocompatibility | journal = Current Topics in Medicinal Chemistry|| volume = 8 | doi= 10.2174/156802608783790901
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