materi biomaterial 2 2

May 2004 Applications of Engineering Mechanics in Medicine, GED – University of Puerto Rico, Mayaguez 1

Learn from yesterday, live for today, dream for tomorrow - - - Chicken Soup for the Soul

BIOMATERIALS FOR ORTHOPEDICS1

Brendamari Rodríguez, Annette Romero, Omar Soto and Oswaldo de Varorna 2

Abstract- Biomaterials deal with the material

aspects of the medical devices. Biomaterials

scientist are concerned with the physical and

chemical properties of materials and their

suitability for a particular device. They are

concerned how these properties are altered by the

biological environment and how the materials may

affect the body. Here we shall discuss selected

biomaterials for orthopaedics.

Keywords- Biomaterials, orthopedics, joints,

artificial joints, 316L stainless steel, Titanium,

Cobalt-Chrome, Zirconium, ceramics, calcium

phosphate, calcium sulfate, materials.

BACKGROUND

Biomaterials improve the quality of life for an ever increasing number of people each year. The range of applications is vast and includes such things as joint and limb replacements, artificial arteries and skin, contact lenses, and dentures. This increasing demand arises from an aging population with higher quality of life expectations. The biomaterials community is producing new and improved implant materials and techniques to meet this demand, but also to aid the treatment of younger patients where the necessary properties are even more demanding. A counter force to this technological push is the increasing level of regulation and the threat of litigation. To meet these conflicting needs it is necessary to have reliable methods of characterization of the material and material/host tissue interactions. The main property required of a biomaterial is that it does not illicit an adverse reaction when placed into service [4].

BIOMATERIALS CLASSIFICATIONS

Biomedical materials can be divided roughly into three main types governed by the tissue response. In broad terms, inert (more strictly, nearly inert) materials illicit no or minimal tissue response. Active materials encourage bonding to surrounding tissue with, for example, new bone growth being stimulated. Degradable, or resorbable __________

1This review article was prepared on May 14, 2004 for the course on Mechanics of Materials – I. Course Instructor: Dr Megh Goyal, Professor in Biomechanical Engineering, Mayaguez Puerto Rico 00681-5984. For details contact: [email protected] or visit at: http://www.ece.uprm.edu/~m_goyal/home.htm 2 The authors are in the alphabetical order. 3The numbers in the parentheses refer to references in the bibliography.

materials are incorporated into the surrounding tissue, or may even dissolve completely over a period of time. Metals are typically inert, ceramics may be inert, active or resorbable and polymers may be inert or resorbable. Table 1 shows examples of biomaterials [4].

Table 1. Types given of biomaterials [4].

Metals Ceramics Polymers

316L stainless steel

Co-Cr Alloys

Titanium

Ti6Al4V

Alumina

Zirconia

Carbon

Hydroxyapatite

Ultra high molecular

weight polyethylene (UHMWPE)

Polyurethane (PE)

APPLICATIONS

The range of applications for biomaterials is large. The number of different biomaterials is also significant. Applications of biomaterials are discussed below:

1. Orthopaedic Applications

Metallic, ceramic and polymeric biomaterials are used in orthopaedic applications. Metallic materials are normally used for load bearing members such as pins and plates and femoral stems etc. Ceramics such as alumina and zirconia are used for wear applications in joint replacements, while hydroxyapatite is used for bone bonding applications to assist implant integration. Polymers such as ultra high molecular weight polyethylene are used as articulating surfaces against ceramic components in joint replacements.

Porous alumina has also been used as a bone spacer to replace large sections of bone which have had to be removed due to disease, [4].

2. Dental Applications

Metallic biomaterials have been used as pins for anchoring tooth implants and as parts of orthodontic devices. Ceramics have found uses as tooth implants including alumina and dental porcelains. Hydroxyapatite has been used for coatings on metallic pins and to fill large bone voids resulting from disease or trauma. Polymers, have are also orthodontic devices such as plates and dentures, [4].


GENERAL REQUIREMENTS

Orthopaedics, like many specialties, has developed through a necessity to correct deformity, restore function and alleviate pain. Orthopaedic surgeons have developed an ability to prevent major losses of bodily function and indeed they can prevent otherwise inevitable death. They seek perfection of their art, by ensuring that the patient reaches optimal condition in the shortest period of time by the safest possible method.

History is very important to an orthopaedic surgeon. The Orthopaedic surgeon has once again been presented with advancing technology. This technology must be applied to the surgeon's practice, but it is best applied only when the surgeon has an underlying knowledge of the history of his art. He must be aware of the way surgeons in the past have contributed to orthopaedics and more importantly, of the mistakes but they have made in the process. The surgeon who makes a mistake that was made by someone before him, is surely humbled and seen as poorly educated. So is he who states that he has developed a technique that no one has thought of before, because chances are that it has been thought of in the past.

In order for orthopaedics to advance in an optimal manner, it is clear that attention must be paid to a history of orthopaedics. The past is our foundation for future developments, we must build upon it so that we too can act as a stable foundation for future generations.

Figure 1. Common sites of infection of bones and joints [16] .

Figure1 shows common sites of infection of bones and joints. It includes pyogenic and tuberculous infection of joints, and osteomyelitis of bones, especially of the hands and feet, and of subcutaneous bones such as the tibiae. In cases when improvement cannot be gained through physical therapy, nonsurgical treatments, or surgical repairs, orthopedic surgeons often advised joint replacement surgery in which the deteriorated joint is removed and replaced with a man-made device. Figure 2 shows a bone plate to assist in the healing of a fracture in the bone. The plate is generally removed once the bone has healed and the bone can support loads without refracturing. Artificial joints consist of a plastic cup made of ultrahigh molecular weight polyethylene (UHMWPE), placed in the joint socket, and a metal (titanium or cobalt chromium alloy) or ceramic (aluminum oxide or zirconium oxide) ball affixed to a metal stem. This type of artificial joint is used to replace hip, knee, shoulder, wrist, finger, or toe joints to restore function that has been impaired as a result of arthritis or other degenerative joint diseases or trauma from sports injuries or other accidents. Joint replacement surgery is performed on an estimated 300,000 patients per year in the U.S. In most cases, it brings welcome relief and mobility after years of pain (Figure 3). After about 10 years of use, these artificial joints often need to be replaced because of wear and fatigue-induced delamination of the polymeric component. Institute engineers are developing improved materials to extend the lifetime of orthopedic implants such as knees and hips


Figure 2. Bone plate, introduced in the early 1900s to assist in the healing of skeletal fractures, were among the earliest successful biomedical implants [5].

Figure 3. Artificial knee joints are implanted in patients with a diseased joint to alleviate pain and restore function [5]. One might think that only surgeons and bioengineers would be involved in improving the design and performance of these implants. Not so. Materials and design engineers (Figure 4) must consider the physiologic loads to be placed on the implants, so they can design for sufficient structural integrity. Material choices also must take into account biocompatibility with surrounding tissues, the environment and corrosion issues, friction and wear of the articulating surfaces, and implant fixation either through osseointegration (the degree to which bone will grow next to or integrate into the implant) or bone cement. In fact, the orthopedic implant community agrees that one of the major

Figure 4. Senior Research Scientist Dr. Cheryl R. Blanchard, Mechanical and Materials Engineering Division of SwRI [5].

problems plaguing these devices is purely materials-related: wear of the polymer cup in total joint replacements. The wear problem plays out a biological disaster in the body. Any use of the joint, such as walking in the case of knees or hips, results in cyclic articulation of the polymer cup against the metal or ceramic ball. Due to significant localized contact stresses at the ball/socket interface, small regions of UHMWPE tend to adhere to the metal or ceramic ball. During the reciprocating motion of normal joint use, fibrils will be drawn from the adherent regions on the polymer surface and break off to form submicrometer-sized wear debris. This adhesive wear mechanism, coupled with fatigue-related delamination of the UHMWPE (most prevalent in knee joints), results in tiny polymer particles being shed into the surrounding synovial fluid and tissues. The biological interaction with small particles in the body then becomes critical. The body’s immune system attempts, unsuccessfully, to digest the wear particles (as it would a bacterium or virus). Enzymes are released that eventually result in the death of adjacent bone cells, or osteolysis. Over time, sufficient bone is resorbed around the implant to cause mechanical loosening, which necessitates a costly and painful implant replacement, or revision. Since the loosening is not caused by an associated infection, it is termed "aseptic”. The average life of a total joint replacement is 8-12 years — even less in more active or younger patients. Because it is necessary to remove some bone surrounding the implant, generally only one revision surgery is possible, thus limiting current orthopedic implant technology to older, less active individuals.

A relatively recent incident in the biomedical device field serves to illustrate the importance of materials choice and engineering on implant performance. The temporomandibular joint (TMJ) provides all jaw mobility and is crucial for chewing, talking and swallowing. This joint can deteriorate from disease or trauma which, in severe cases, necessitates replacement by an artificial joint. For many years, less than optimum technologies existed for TMJ implants. In the late 1970s, a TMJ replacement using polytetraflouroethylene (PTFE) as the bearing counterface was invented, and, in 1983, the inventors received FDA approval to market the PTFE implant, which was called the


Interpositional Implant (IPI). In theory, PTFE would seem an appropriate choice for an implant material, as it exhibits a low coefficient of friction and has been used extensively as a bearing surface in other engineering applications. However, of the more than 25,000 PTFE TMJ implants received by patients, failed. The fibrillation and small particles are characteristic of an adhesive wear mechanism, which

can result in surrounding bone loss and the need for implant replacement (Figure 5).

Figure 5. These micrographs, taken at a magnification of 20,000X on a scanning electron microscope, illustrate the wear problem that occurs with an artificial joint implant component (socket) constructed of UHMWPE. At top is unworn UHMWPE. The UHMWPE sample has undergone a friction and wear test versus cobalt chromium (artificial joint ball material) [5].

A low coefficient of friction of PTFE is due to formation of a thin film of the material onto the opposing bearing surface. Although this transfer film acts as a lubricant, it also, by virtue of its formation, subjects the material to an adhesive wear mechanism. In the case of the PTFE TMJ implants, surrounding tissues quickly became overwhelmed by wear debris, and the immune system response result in osteolysis, causing massive destruction of the joint and surrounding tissues. For those people who received the implants, this was truly a tragedy; many suffered severe facial deformities, and most experienced unbearable pain and were no longer able to chew, swallow, or sleep. At the time the IPI was developed, evidence did exist that PTFE was not an appropriate implant material.

In the late 1950s, Dr. John Charnley, at Wrightington Hospital in the U.K. pioneered the first total hip replacements using PTFE as the cup bearing surface. Dr. Charnley reported massive wear of the PTFE part and early clinical failure as a result of aseptic loosening. These findings, reported widely in the open literature and in later reports from researchers testing the IPI implant, should have been sufficient warning that PTFE was not an appropriate material to use as a load-bearing surface in the body. Work at SwRI is addressing the wear problem in UHMWPE total joint prostheses.

In collaboration with scientists at the University of Texas Health Science Center at San Antonio funded

by the National Science Foundation, SwRI scientists and engineers are studying the wear process and biological responses to wear debris. Results of these studies have led to novel ideas for materials modification and development. The Institute is also developing new composite materials to defeat the fatigue-induced delamination observed in the UHMWPE component of knee implants. Studies of wear debris extracted from actual tissue samples of patients whose implants failed as a result of aseptic loosening have generated significant information regarding wear particle size, shape, and surface morphology. Institute scientists were the first to use the atomic force microscope (AFM) to produce detailed, high resolution images of wear particles. A few hundred nanometers in size, the UHMWPE wear debris studied at SwRI sometimes exhibits a cauliflower-like surface morphology. Scientists at the Health Science Center will use similar particles to study the biological response elicited by the particles. By combining wear debris and cellular response studies, engineers and biologists will be able to better understand implant failure and to re-engineer implants to prevent future problems, [ 5, 16 ].

BONE GRAFT SUBSTITUTES

In many cases, the loss of bone due to surgery, accidents or normal aging requires the substitution of bone in order to facilitate the rehabilitation of the patient. Figure 6 shows two cases in which bone substitution is required.

a. Collapsed disc. b. Non-Union

Figure 6. Examples where bone substitution is required [24]

Nowadays, the need for bone substitutes includes autograftings procedures, allograftings procedures or synthetic bone substitutes. Autografting, which represents about 58% of the current bone substitutes, involves harvesting a bone from one location in the patient’s body, usually taken from the pelvic region, and transplanting it into another part of the same patient. Using this procedure, when the autogenous grafts are available, typically produces the best clinical results. This procedure has obvious benefits, like the elimination of immunogenicity problems. Autografting, however, has several associated problems including the additional surgical costs for the harvesting procedure, infection, pain at the harvesting site and that the sample of the patient’s own bone that can be taken is very small, among other things. The allografting procedure consists in harvesting and processing bone from a live or deceased donor and then transplanting it to the patient. These implants are acellular and are less successful than autografts implants for reasons attributed to


immuogenicity and the absence of viable cells that become osteoblasts. Another disadvantage of allografting is relate with transmitted disease. Due to complications related to these procedures, bone graft substitutes made with synthetics materials are becoming very important in bone substitutions procedures. The ideal bone graft substitute should be osteogenic, biocompatible, bioabsorbable, able to provide structural support, easy to use clinically, and cost-effective. The bone grafts and their substitutes can be divided according to their properties of osteoconduction, osteoinduction, and osteogenesis. Table 2 shows classifications of bone graft substitutes. The synthetic material belongs to the osteoconductive category. The osteoconductive synthetic grafts that are used generally falls under the calcium sulphate and calcium phosphate groups. They can be used as preset and injectable materials. The next sections will present a more detailed information about synthetic bone graft, calcium sulphate and calcium phosphate materials, [6, 7]. Table 2. Classification of Bone Graft Substitutes Based on Properties [6].

Description Classes

Osteoconduction Provides a passive porous scaffold to support of direct bone formation

Calcium sulphate, ceramics, calcium phosphate cements, collagen, bioactive, glass, synthetic polymers

Osteoinduction Induces a differentiation of stem cells into osteogenic cells

Demineralized bone matrix, bone morphogenic, proteins, growth factors, gene therapy

Osteogenesis Provides stem cells with osteogenic potential, which directly lays down new bone

Bone marrow aspirate

Combined Provides more than one of the above

Composites

1. Indications for Bone Substitutes The main indications for bone substitutes will be in spinal fusion, bone defects, osteoporotic fractures, revision surgery and, recently, vertebroplasty (injecting a vertebra with synthetic material). Vertebroplasty using polymethylmethacrylate was first introduced in France more than 15 years ago by neurosurgeons, but its use is now spreading rapidly.

This mini-invasive procedure for the treatment of vertebral fractures in osteoporosis can reinforce fractured bone, alleviate chronic pain and prevent further vertebral collapse. Vertebroplasty is performed under biplanar fluoroscopic control, CT or guided navigation (Figure 7) [6, 7].

Figure 7. Schematic drawing of a vertebral body, on which synthetic bone grafting (vertebroplasty) is performed using the injection-suction method. Two needles are used, one for injecting the synthetic bone material and the other for developing an underpressure in the vertebral body. This method reduces the risk of leakage into vessels or the nerves in the spinal canal, [7].

Figure 8. Fracture treated with synthetic bone graft and internal fixation [7].

2. Synthetic Bone Grafts

Synthetic materials can be made from different materials that are biologically compatible materials. The synthetics materials exhibits the property of being osteoconductives materials, which mean being bone-stimulating materials. Some of these materials can be mixed with bone marrow aspirate to obtain osteoinductive properties (bone forming). Some important characteristics of synthetic bone substitutes are:

- Its porosity (determines the amount of surface

area expose to bone tissue ingrowth). Porosity alone is not adequate for bone ingrowth.


- Resorption rate (ability to disappear so as to be replaced by new bone).

- Biocompatibility to prevent inflammatory

reactions, minimizing the interference with bone induction.

- Biodegradable so that the patient’s own

bone can replace the foreign substitute.

The osteoconductive synthetic grafts that are used generally are made of calcium sulphate and calcium phosphate materials. Hydroxyapatite synthetic bone , a derivate of calcium phosphate, is an important material due to its biocompatibility. Figure 8 shows an example of a fracture treated with injectable synthetic bone and internal fixation [1, 6].

TITANIUM AND TITANIUM ALLOYS

Although Titanium has excellent heat and corrosion resistance capabilities, it is extremely difficult to form and machine into desired shapes. Also its extreme chemical reactivity with air, combined with other factors, has caused the cost of titanium components to be very high. The only economical applications of this material currently (until more efficient techniques of working with it can be found) are in aerospace applications where weight and temperature resistance are very important, and in military applications, they provide extreme corrosion resistance and durability. Titanium is also used in biomedical applications such as prosthetics and implants ( Figure 9) due to its biological inertness. There are several titanium alloys that have been developed for use in the past four decades. These alloys include Ti-6Al-4V (an alloy of titanium, aluminum and vanadium), the most highly used alloy of titanium and Ti-4Al-4Mo-2Sn-0.5Si (an alloy of titanium, aluminum, molybdenum, tin, and silicon), which was developed later and is used less frequently. Table 3 presents some properties of the titanium alloy Ti-6Al-4V [17].

Figure 9. Example of a titanium biomedical applications [17].

Table 3. Properties of Ti-6Al-4V at 25°C [22].

Property of the Ti-6Al-4V Values of the Ti-6Al-4V

Density 4430

Poisson's Ratio 0.34

Elastic Modulus GPa

113.8

Tensile Strength MPa

993

Yield Strength MPa

924

Elongation %

14

Reduction in Area %

30

Hardness HRC

36

a. Physiological Behavior

These materials are classified as biologically inert biomaterials or bioinert. As such, they remain essentially unchanged when implanted into human bodies. The human body is able to recognize these materials as foreign, and tries to isolate them by encasing them in fibrous tissues. However, they do not illicit any adverse reactions and are tolerated well by the human body. Furthermore, they do not induce allergic reactions such as has been observed on occasion with some stainless steels, which have induced nickel hypersensitivity in surrounding tissues. The surface of titanium is often modified by coating it with hydroxyapatite. Plasma spraying is the only commercially accepted technique for depositing such coatings. The hydroxyapatite provides a bioactive surface (i.e. it actively participates in bone bonding), such that bone cements and other mechanical fixation devices are often not required [18].

b. Mechanical Suitability

Titanium and its alloys possess suitable mechanical properties such as strength, bend strength and fatigue resistance to be used in orthopaedics and dental applications. This is why they have been employed in load-bearing biomedical applications instead of materials such as hydroxyapatite, which displays bioactive behavior. Other specific properties that make it a desirable biomaterial are density and elastic modulus. In terms of density, it has a significantly lower density (Table 4) than other metallic biomaterials, implying that the implants will be lighter than similar items fabricated out of stainless steel or cobalt chrome alloysHaving a lower elastic modulus compared to the other metals is desirable as the metal tends to behave a little bit more like bone itself, which is desirable from a biomechanical perspective. This implies that the bone hosting the biomaterial is less likely to atrophy and resorb [18].


Table 4: Density and elastic modulus of selected biomaterials [18].

Material Density Elastic

Modulus

Cortical Bone ~2.0 g.cm-3 7-30 GPa

Cobalt-Chrome alloy ~8.5 g.cm-3 230 GPa

316L Stainless Steel 8.0 g.cm-3 200 GPa

CP Titanium 4.51 g.cm-3 110 GPa

Ti6Al4V 4.40 g.cm-3 106 GPa

Figure 10. Implant components for a total hip replacement (photo courtesy of Dr. Karlis Gross) [18].

c. Applications

Titanium is commonly used in orthopaedic implants such as joint replacements and bone pins, plates and screws. Figure 10 shows the various components of a total hip replacement. On the left is the femoral stem made of a titanium alloy. The long round section fits down into the thigh bone or femur. The white section is a hydroxyapatite coating to encourage bone bonding to the implant. This section is also macrotextured to provide surface features for the bone to mechanically interlock with. The ball on top of the femoral stem is called the femoral head. It is made of zirconia ceramic and fits into the hip joint in the pelvis.

The hemispherical item on the right is the acetabular cup, also made from titanium alloy. It is coated with porous alumina ceramic, to allow bone ingrowth for stabilisation. A ultra high molecular weight polyethylene (UHMWPE) liner fits inside the acetabular cup and provides the articulating surface for the femoral head.

Figure 11 shows a prototype total knee replacement prosthesis, similar in design to many commercial implants. It consists of titanium alloy upper and lower structural components. A zirconia wear surface has been fabricated for the upper section. Similar to the hip prosthesis, this articulates against a UHMWPE insert on the lower section. Other orthopaedic applications for titanium-based materials include bone pins, plates and screws, used for repairing broken bones etc [18].

Figure 11. A total knee replacement prosthesis (photo courtesy of Dr. Besim Ben-Nissan) [18].

CERAMICS

For many years, ceramic materials were only useful in the making of pottery and other artwork. They have since evolved into one of the most important biomaterials used today because of their beneficial properties. Ceramic materials are nonmetallic, inorganic compounds that exhibit great strength and stiffness, resistance to corrosion and wear, and low density. These characteristics allow ceramics to become prime candidates for a wide range of biomedical applications. Ceramics are used in several different fields such as dentistry, orthopaedics, and as medical sensors. In dentistry, ceramics are commonly used for implants such as crowns and dentures. The orthopedic field utilizes ceramics for joint and bone segment replacement and temporary bone repair devices. Ceramics are also used as coatings for implants made of other materials to provide a biocompatible interface with the body [9].

1. Bone graft substitutes: Calcium phosphate

The first applications of calcium phosphate salts were powders. The ceramic form first became available in the 1960s and was later evaluated as a bone graft substitute. The synthetic hydroxyapatite is one of the most commonly used calcium phosphate ceramics. Synthetic ceramics provide an osteoconductive scaffold to which chemotactic, circulating proteins and cells can migrate and adhere, and within which progenitor cells can differentiate into functioning osteoblasts. Ceramics do not supply osteogenic cells as found in autograft. They do not have even the weak osteoinductive potential found with allograft. However, ceramics are readily available and bypass the known risks of allograft-induced immunogenic response or disease conveyance, as well as surgical complications from retrieving bone from an autogenous second site. The chemistry, architecture, shape, and positioning of the ceramic material influence the speed and extent of remodeling. Its bioresorbability depends on the amount of surface area exposed, which is governed, in turn, by crystal size, the form supplied, and density. A ceramic material formed as a dense block exposes only a small surface area, thus slowing or confining surfaces accessible for


resorption. Most calcium phosphates are classified as resorbable biomaterials. This means that under physiological conditions they will dissolve. Table 5 shows examples of calcium phosphate compounds. The benefit of calcium phosphate biomaterials is that the dissolution products can be readily assimilated by the human body. Calcium phosphate is mainly used in filling defects (for example, areas of bone loss such as in tibial plateau fracture), in composite grafts to supplement autograft, and at sites where compression (rather than tension, bending, or torsion) is the dominant mode of mechanical loading. Variation in the properties of calcium phosphate coatings has an effect on the bone-bonding mechanism and the rate of bone formation. Both the composition and the crystallinity of the calcium phosphate coating are important parameters that determine its bioactivity characteristics. Hydroxyapatite {Ca10[PO4]6[OH]2} has the ability to bond to osseous and epithelial tissue. Hydroxyapatite, with beneficial bone tissue growth effects, is used as a coating material since it does not have sufficient strength and toughness to be used by itself as biomedical implants. Unlike other calcium phosphates, hydroxyapatite does not break down under physiological conditions. Under normal physiological conditions of pH 7.2, hydroxyapatite is the stable calcium phosphate compound. This may drop to as low as pH 5.5 in the region of tissue damage, although this would eventually return to pH 7.2 over a period of time. Even under these conditions hydroxyapatite is still the stable phase. It actively takes part in bone bonding, forming strong chemical bonds with surrounding bone. This property has been exploited for rapid bone repair after major trauma or surgery. Figure 12 shows bone ingrowth around synthetic hydroxyapatite. While its mechanical properties have been found to be unsuitable for load-bearing applications, it is used as a coating on materials such as titanium and titanium alloys, where it can contribute its 'bioactive' properties, while the metallic component bears the load. Such coatings are applied by plasma spraying. However, careful control of processing parameters is necessary to prevent thermal decomposition of hydroxyapatite into other soluble calcium phosphates due to the high processing temperatures [3, 6].

2. Calcium Sulphate (Plaster of Paris)

Although its external use for creation of hard setting bandages dates back to the seventeenth century, the first internal use of Gypsum (Plaster of Paris) to fill bony defects was reported in 1892 by Dressmann. The application of Plaster of Paris as a bone void filler, and the use of antibiotic-laden plaster in the treatment of infected bony defects has been supported by various studies. Peltier reported the initial results of its use and recommended it as a cheap, bioabsorbable bone graft substitute. Calcium sulphate

Table 5. Calcium phosphate compounds as biomaterials [3].

Chemical

Name

Abbr Chemical

Formula

Phase Ca/P

Amorphous calcium

phosphate ACP - - -

Dicalcium Phosphate

DCP CaHPO4 Monetite 1.00

Tricalcium Phosphate

α-TCP Ca3(PO4)2 1.50

Tricalcium Phosphate

β-TCP Ca3(PO4)2 Whitlock

ite 1.50

Pentacalcium Hydroxyl Apatite

HAp Ca10(PO4)6(O

H)2 Hydroxyapatite

1.67

Tetracalcium Phosphate Monoxide

TTCP Ca4O(PO4)2 Hilgenstockite

2.00

Figure 12. Bone ingrowth around synthetic hydroxyapatite [7].

(CaSO4) has long been used in its partially hydrated form. When mixed with water, it initiates an exothermic reaction that leads to recrystallization of the calcium sulfate into the solid form. When it is mixed with The problem with this reaction is that the recrystallization proceeds randomly, producing crystals of varying size and shape as well as multiple defects within the structure. This variability in the crystalline structure causes significant variability in solubility, mechanical properties, and porosity. In addition, it may resorb too rapidly, leading to fibrous ingrowth instead of bony substitution. Medical grade calcium sulfate is crystallized in highly controlled environments producing regularly shaped crystals of similar size and shape. It possesses a slower, more predictable solubility and resorption. One such material is OsteoSet (Wright Medical Technology, Arlington, Tenn), which was approved by the FDA in 1996. The material is available in 3- and 4.8-mm pellets that typically dissolve in vivo within 30-60 days, depending on the volume and location. The pellets are packaged in vials and are sterilized by gamma irradiation. It also is available in a powdered form, the OsteoSet Resorbable Bead Kit (Wright Medical


Technology, Arlington, Tenn), standard or fast cure (5 minutes setting time for fast cure kit compared to 20 minutes for the standard kit), thus maximizing the surgical options for adding antibiotics and filling defects with custom molded beads or shapes. The chief advantage is that it can be used in the presence of infection. Because it is bioabsorbable, it has inherent advantages over other antibiotic carriers, such as polymethylmethacrylate, which become a nidus for further infection after elusion of the antibiotics, thus requiring a separate operation for removal from the surgical site. When combined with the eradication of dead space and the acidic environment created during its resorption, the compound can be an effective treatment for acute bony infections with bone loss. The disadvantages of calcium sulphates are their weak mechanical strength and rapid resorption within 6–12 weeks. For clinical use, injectable osteoconductive grafts should ideally be biphasic with a compressive strength >25 Mpa. Their injection time should be between 2 and 6 min, with a setting time of less than 10 min [3, 6].

3. Disadvantages of Ceramics

A shortcoming noted with ceramics used as stand-alone bone substitutes is the initial low resistance to impact and fracture. Due to its brittle structure, use of a ceramic material in conditions of torsional, impact, or shear stress is limited. However, cancellous bone grafts likewise contribute little immediate structural support prior to union with the host site and remodeling along lines of stress. Another disadvantage found with ceramic implants is the difficulty of radiographic assessment of the ingrowth into the defect site until partial resorption has occurred [1].

COBALT AND COBALT CHROME

Cobalt and Cobalt Chrome

Cobalt

Brandt discovered cobalt around 1735. It occurs in the minerals cobaltite, smaltite and erythrite and is often associated with nickel, silver, lead, copper and iron ores, from which it is most frequently obtained as a by-product. It is also present in meteorites. Cobalt is a brittle, hard metal white in appearance resembling nickel (and iron) but with a bluish tinge instead of the yellow of nickel. It is rarer and more valuable than nickel. It is diamagnetic and has magnetic permeability approximately two thirds that of iron and three times that of nickel. Cobalt exists as two

allotropes over a wide temperature range. The β-form a close-packed hexagonal crystal is stable and predominates below approximately 417°C (782°F),

and the α-form a cubic crystal is stable and predominates above this temperature until the melting point. Although allied to nickel, it is more active

chemically than nickel. Cobalt dissolves in dilute sulphuric acid, nitric or hydrochloric acid and is slowly attacked by alkalis. The oxidation rate of pure cobalt is twenty five times that of nickel. Cobalt’s ability as a whitening agent against copper alloys is inferior to that of nickel. However, small amounts in nickel-copper alloys will neutralise the yellowish tinge of the nickel and make them whiter. Cobalt imparts red-hardness to tool steels. It can harden alloys to greater extent than nickel, especially in the presence of carbon and can form more chemical compounds in alloys than nickel. Natural cobalt is cobalt 59, which is stable and non-radioactive, but other isotopes 54 to 64 are all radioactive (table 6), emitting beta and gamma radiation. Other isotopes not listed in table 1 have short half-lives.[3]

Table 6. Cobalt isotopes and their half-lives. [3]

Isotope Half Life

Cobalt 60

Cobalt 58

Cobalt 57

Cobalt 56

5.3 years

72 days

270 days

80 days

Aplications:

Cobalt 60 has a number of applications. These include:

· Radiographic inspection · A gamma ray source · A tracer · A radiotherapeutic agent · Irradiation of plastics · A catalyst for the sulphonation of paraffin oil. In this

application the gamma rays emitted by the cobalt cause the reaction of sulphur dioxide and liquid paraffin.

Other uses for cobalt are:

· In superalloys for aircraft gas turbine engines · It is a key elemental ingredient in magnet steels, by

which it increases residual magnetism and coercive force and in nonferrous-base magnetic alloys

· Cobalt is an important element in numerous glass-to-metal sealing alloys as well as low expansion alloys

· Alloys for dental and surgical applications because they are not attacked by physiological fluids. An example of which is ‘Vitallium’ which is used to replace bone. Such alloys are ductile enough to permit anchoring of dentures on neighbouring teeth and contain up to 65% cobalt

· High-speed, heavy duty, high temperature cutting tools, and dies

· Gas turbine generators · Electroplating.


Cobalt salts are used as a source of brilliant permanent blue colour in porcelain, glass, pottery, tiles and enamels [22, 23].

Properties

Creep Resistance

One of the main attractions of cobalt-based alloys is their excellent creep resistance. Materials creep due to thermally-activated movement of dislocations through a crystalline matrix. These alloys possess a matrix that is resistant to this as cobalt has a good tolerance for other elements in solid solution. These elements can effectively strengthen the matrix. Their ability to do this depends on factors such as:

· The difference in atomic size between cobalt and the solute

· The effect of the solute on the stacking fault energy

· The diffusion rate of the solute into the cobalt matrix

It has also been found that a matrix containing a larger number of solutes is often better than one containing a fewer number, hence the strengthening of the matrix is also dependent on the amount of alloying elements available to the go into solid solution, that have not formed carbides, or for that matter intermetallics. The key elements for this process are chromium, tungsten, niobium and tantalum. A second strengthening mechanism also exists and involves the formation of carbides and carbonitrides forming with chromium (primarily), tungsten, molybdenum, niobium, tantalum, zirconium, vanadium and titanium. Carbides formed include MC, M6C, M7C3, M23C6 and sometimes M2C3, with the amount of each depending on factors such as availability of elements to form carbides, carbon content and thermal history. It is also possible for nitrogen to substitute for carbon in these structures. Optimum properties are produced when carbides precipitate both intergranularly and intragranularly. Intergranular precipitation prevents gross sliding and grain boundary migration and can form a skeleton if present in sufficient quantities, while intragranular precipitation strengthens the matrix by inhibiting the motion of dislocations. Carbide distribution by solidifaction parameters such as pouring temperature and cooling rate. As cast alloys are rarely heat treated, carbides will generally only form during prolonged exposure to operating temperatures. Wrought materials on the other hand may be hot worked. Further strengthening can be induced by solution heat treatment between 1175-1230°C and rapid cooling [22, 23].

Room Temperature Properties

RoomAs these alloys are generally used at elevated temperatures, the room temperature properties are not

relevant to the service conditions. They do however, play a role for manufacturers e.g. tensile strength and ductility can influence how much hot or cold working the material can withstand and hardness influences machinablity. It should also be noted that room temperature properties such as elongation can be effected by the thermal history of the material, i.e. amount of carbide precipitation, with more precipitation leading to lower ductility. Also increased exposure to high temperatures increases the hardness of higher carbon alloys more so than lower carbon content

alloys.

It has been shown that cobalt-chrome alloys with veneering capacity, such as Wirobond®C, represent an alternative to alloys with a high gold content . As far as corrosion and biocompatibility are concerned, both groups can be designated as equivalent. In mechanical terms (modulus of elasticity, heat resistance, thermal conductivity), cobalt chrome is superior to alloys with a high gold content. In terms of price, the cobalt chrome alloys again have an advantage. Both alloys and alloy types display adequately high shear bond strength. The experimental results show higher values for the gold alloys, though whether this is of clinical relevance is still a matter of debate [22, 23]. Properties

• Cobalt 61.0% base metal

• Casting temperature Higher, therefore cannot be cast with all casting machines

• Investment materials No investment materials containing plaster may be used because the required preheating temperature would lead to decomposition of the investment material. The decomposition products resulting from this react strongly to the alloy melt flowing in. No investment materials containing graphite may be used since chromium carbide would otherwise be formed, leading to extreme hardening (HV 10 > 700).

• Finishing Increased work requirement and greater wear of the equipment (grinding stones, milling units, etc.). Consequently the price advantage (Wirobond C: approx. DM 0.82/g; Bio PontoStar: approx. DM 33.00/g) is reduced, but not eliminated.

• Tensile strength Higher than gold alloys

• Modulus of elasticity Approximately double, which means that Wirobond®C has a significantly higher load capacity with equal modelling strength. This is of interest for long-term stability.

• 0.2% ductile yield Comparable, measure for the permanent deformation (important with clasps)

• Elongation limit Lower, not of crucial importance for the finishing capacity for crown and bridge alloys

• Hardness Higher, difficult to finish

• Coefficient of thermal expansion Comparable, see also veneering capacity

• Heat resistance Significantly higher, particularly in comparison to palladium-free gold alloys, thus


more secure against distortion during veneering process

• Thermal conductivity Lower than with gold alloys, therefore greater wearing comfort for the patient [3, 22].

Cobalt vs. Stainless Steel 316L

Stainless Sell 316L

Grade 316 is the standard molybdenum-bearing grade, second in importance to 304 amongst the austenitic stainless steels. The molybdenum gives 316 better overall corrosion resistant properties than Grade 304, particularly higher resistance to pitting and crevice corrosion in chloride environments. It has excellent forming and welding characteristics. It is readily brake or roll formed into a variety of parts for applications in the industrial, architectural, and transportation fields. Grade 316 also has outstanding welding characteristics. Post-weld annealing is not required when welding thin sections.

Grade 316L, the low carbon version of 316 and is immune from sensitisation (grain boundary carbide precipitation). Thus it is extensively used in heavy gauge welded components (over about 6mm). Grade 316H, with its higher carbon content has application at elevated temperatures, as does stabilised grade 316Ti.

Characterised by high corrosion resistance in marine and industrial atmospheres, it exhibits excellent resistance to chloride attack and against complex suphur compounds employed in the pulp and paper processing industries. The addition of 2% to 3% of molybdenum increases its resistance to pitting corrosion and improves its creep resistance at elevated temperatures. The low carbon content reduces the risk of intergranural corrosion (Due to carbide precipitation) during welding, reducing the need for post weld annealing. Finally it displays good oxidation resistance at elevated temperatures.

Stainless steel 316L cannot be hardened by thermal treatment, but strength and hardness can be increased substantially by cold working, with susequent reduction in ductility.

It is now available with improved machinability (by calcium injection treatment), which has little effect on corrosion resistance and weldability while greatly increasing feeds and/or speeds, plus extending tool life.

Typical uses are: Architectural Components, Textile Equipment, Pulp and Paper Processing Equipment, Marine Equipment and Fittings, Photographic Equipment and X-Ray Equipment etc..

Material non magnetic in the annealed condition, but can become mildly magnetic following heavy cold working. Annealing is required to rectify if necessary [3, 22, 23].

Mechanical Properties

Table 7. Mechanical properties of 316 grade stainless steels.[3, 23]

Hardness Grade Tensile Str

(MPa) min

Yield Str

0.2% Proof (MPa) min

Elong (% in

50mm) min

Rockwell B (HRB) max

Brinell (HB) max

316 515 205 40 95 217

316L 485 170 40 95 217

316H 515 205 40 95 217

Note: 316H also has a requirement for a grain size of ASTM no. 7 or coarser.

Physical Properties

Table 8. Typical physical properties for 316 grade stainless steels. [3, 23].

Mean Co-eff of Thermal Expansion (µm/m/°C)

Grade Density (kg/m3)

Elastic Modulus

(GPa) 0-100°C 0-315°C

0-538°C

316/L/H 8000 193 15.9 16.2 17.5

Table 9. Possible alternative grades to 316 stainless steel [3].

Grade Why it might be chosen instead of 316?

316Ti Better resistance to temperatures of around 600-900°C is needed.

316N Higher strength than standard 316.

317L Higher resistance to chlorides than 316L, but with similar resistance to stress corrosion cracking.

904L Much higher resistance to chlorides at elevated temperatures, with good formability

2205 Much higher resistance to chlorides at elevated temperatures, and higher strength than 316

Corrosion Resistance

Excellent in a range of atmospheric environments and many corrosive media - generally more resistant than 304. Subject to pitting and crevice corrosion in warm chloride environments, and to stress corrosion cracking above about 60°C. Considered resistant to potable water with up to about 1000mg/L chlorides at ambient temperatures, reducing to about 500mg/L at 60°C.

Stainless Steel 316 is usually regarded as the standard “marine grade stainless steel”, but it is not resistant to warm sea water. In many marine environments 316 does


exhibit surface corrosion, usually visible as brown staining. This is particularly associated with crevices and rough surface finish [22].

Heat Resistance and Temperature Properties

Good oxidation resistance in intermittent service to 870°C and in continuous service to 925°C. Continuous use of 316 in the 425-860°C range is not recommended if subsequent aqueous corrosion resistance is important. Grade 316L is more resistant to carbide precipitation and can be used in the above temperature range. Grade 316H has higher strength at elevated temperatures and is sometimes used for structural and pressure-containing applications at temperatures above about 500°C.

316L displays good oxidation resistance in continuous service up to 930 oC, and in intermittent service up to 870 oC. Due to its low carbon content it is also less susceptable to carbide precipitation resulting in intergranular corrosion when heated or slow cooled through the temperature range 430 oC - 870 oC either in service or during welding. There is however a reduction in mechanical properties as temperature increases [22].

Applications

Typical applications include:

· Food preparation equipment particularly in chloride environments.

· Laboratory benches & equipment.

· Coastal architectural panelling, railings & trim.

· Boat fittings.

· Chemical containers, including for transport.

· Heat Exchangers.

· Woven or welded screens for mining, quarrying & water filtration.

· Threaded fasteners.

· Springs [22, 23].

Thermal Properties

Thermal expansion properties are similar to those of nickel-based alloys.

Thermal conductivity values for cobalt-based carbide-hardened alloys such as HS 21 are typically about 15% of those for pure cobalt [22, 23] .

Oxidation Resistance

This property is almost entirely dictated by the chromium content. Chromium contents in the range 20-25% are usually sufficient to protect the alloy up to temperatures of 1100°C. Although the chromium is responsible for the formation of a protective oxide layer, it is susceptible to attack from elements such as sulphur, vanadium and alkali metal halides or oxides. These commonly come from contaminatyed fuels and other sources. Sulphur penetration caqn lead to the formation of sulphides within the alloy, forming low melting point eutectics such as Co4S3 (melting point 877°C). Strengthening carbides may also be preferentially attacked in some alloys [22, 23].

Cobalt Chrome:

Cobalt-chrome alloys are part of the group of non-precious alloys, also referred to as preciousmetal- free alloys. The first cobalt-chrome alloy that was introduced in dentistry in the1930s was an alloy used in medical implantology, where it had already proven its clinicaleffectiveness. It was used in the partial denture technique and replaced steel in that field. In dental usage the term “steel” became a synonym with partial denture alloys consisting ofcobalt-chrome alloys. However, this designation is misleading since steel refers to iron alloyscontaining carbon. The frequently used designation “chrome-cobalt alloy” is also incorrectbecause by definition this would involve alloys on a chromium base. What are meant arealloys on a cobalt base. Besides being used as partial denture alloys, cobalt-chrome alloys such as Wirobond®C (BEGO) can be utilised as crown and bridge alloys for ceramic veneering. The acrylic veneering of cobalt-chrome alloys generally displays more favourable bond values than with precious-metal alloys. Non-precious alloys have a negative reputation among some dentists and dental technicians. Poor processability, inadequate chemical and biological properties are cited as reasons for this. This polarisation goes so far that consideration is only given to alloys with a high gold content, whose properties are applied to other precious-metal alloys (on palladium or silver base and to alloys with reduced gold content) without reflection, however. This results in a distorted picture that does not accurately reflect the non-precious alloys [3, 22, 23]. . Composition

A carbon content of less than 0.02% ensures that no carbide precipitation that would lead to brittleness of the marginal areas of the seam occurs during laser welding. This would then result in an increased risk of fracture. Highly pure base metals are used to make alloys. However, there are no 100% pure metals. For example, platinum ores contain palladium and sometimes also nickel impurities, cobalt is accompanied by nickel (and conversely), etc. Complete


separation of the elements is never possible. The relevant standards stipulate a maximum nickel content of 0.1 %. Concentrations of greater than 0.1% have to be declared. Alloys with less than 0.1% of nickelcan be designated as nickel-free. The claim that a cobalt-chrome alloy is absolutely nickel-free would be objectively false and can only be understood on the basis of marketing aspects. If a restoration made of a cobalt-chrome alloy weighed 10 g (extremely large bridge, partial denture), the entire restoration would contain a maximum of 0.07 g (= 70 mg) of nickel. The latter, however, is not only found on the alloy surface, but is spread homogeneously throughout the restoration. If one assumes that nickel is detached from the alloy to the same extent as cobalt (which is probable although nickel is nobler than cobalt), the release of nickel will amount to approx. 0.00003 mg/cm² (0.03 µg/cm²) in the first week and constantly decline thereafter. If one compares this to the daily uptake in food, i.e. approx. 0.19 – 0.90 mg, (190 –900 µg), toxicological or allergic stress appears very improbable. In the case of alloys with veneering capacity, the available area is additionally reduced considerably due to the veneered ceramics [3].

Dental processing and mechanical values

Can be used for veneering crowns and bridges with ceramics. The dental processing of cobalt-chrome alloys is assessed as more unfavourable in comparison to gold alloys. This is also reflected in the slightly higher costs for the required instruments. This partially offsets the price advantage of the alloy. This opinion must be qualified, however. In the veneering of frames the difference in the required processing between gold and cobalt-chrome alloys with veneering capacity is not very great. In the case of fully cast crowns, the more difficult processing of the cobalt-chrome alloys is a significant negative factor. It is recommended, therefore, that the processing instructions be followed. Each alloy has its specific features that must be taken into account. This applies to non-precious alloys as well as to precious-metal alloys [3, 22, 23]. .

Veneerability

Due to the higher melting interval, non-precious alloys are generally more heat-resistant than gold alloys. In particular palladium-free gold alloys are sensitive here since palladium is responsible for the heat resistance, among other things. Heat resistance refers to the ability of an alloy not to deform even in the high-temperature range (slightly below the solidus point), i.e. not to distort under its own weight. With a coefficient of thermal expansion of 14.2 [10 –6 * K.

Corrosion

As already explained in connection with the composition of the alloy, chromium and molybdenum are important for corrosion resistance. The latter can be tested with an immersion test. Test objects are suspended in a solution consisting of sodium chloride and lactic acid (0.1 mol/l each) and the dissolved alloy components are determined by means of a suitable analytical method (e.g. atomic absorption spectrometry, AAS). The ion quantities determined can then be compared to other alloys By comparing the corrosion rates of comparable and clinically proven alloys, conclusions can be made concerning the behaviour of the alloy examined. This study method is therefore suitable as a pre-clinical screening test. It has been shown that cobalt-chrome alloys display an ion release that is somewhat higher than that of gold alloys, but is still on the same order of magnitude. It is known that dental processing, such as casting, grinding or ceramic veneering, may influence the corrosion characteristics of dental alloys. In the case of cobalt-chrome alloys, this influence are relatively small. This means that such alloys are very rugged [3, 22, 23] .

Biocompatibility

The main components of cobalt-chrome alloys, cobalt, chromium and molybdenum, are essential elements]. Therefore, they must be classified as more favorable in principle, as elements that have no function in the human body. For essential elements the human organism has diverse ways of decomposition and utilization. There appear to be certain threshold values, below which no interaction takes place. However, these threshold values are very individual and may be very low in specific cases since there is verification of allergies to cobalt, chromium and molybdenum. In the relevant literature, however, there is no reference to the fact that cobalt-chrome alloys have caused an allergy. This point is also supported by the use of cobalt-chrome alloys for partial dentures for decades. Alloys of this type were positively assessed back in 1936. Thus, there is no clinical experience in this connection that is older than that concerning gold alloys with veneering capacity. Allergies are usually verified by means of the patch test. It must be emphasised here that this test itself is capable of sensitising the subject. Therefore, in Norway, for example, it is only permitted if there is justified suspicion of an allergy. Furthermore, there is problem regarding suitable test substances. For some elements there are still none, with others the selection of unsuitable test substances may lead to incorrect statements. Chromium allergies (for the dental field), for example, should only be tested with chromates, in which chromium is found with the oxidation number +III. If one uses dichromate (here chromium has the oxidation number +VI and acts as a strong oxidant), one will obtain in all likelihood incorrect results. Chromium with the oxidation number +III is released from dental alloys due to corrosion processes. To avoid faulty diagnoses, the patch test should only be conducted by properly trained persons (dermatologists, allergists)


because the evaluation requires know-how and experience [3, 22, 23] .

Table 10. Biomaterials – Densities of Biomaterials and Some Other Related Materials [3].

Material Density

(g/cm3)

Material Density

(g/cm3)

Amalgam 11.6 Palladium based alloys

~10.8

Alumina 3.85 Porcelain dental

~2.05

Bone - cortical ~2.0 Stainless steel – 316L

8.0

Calcium hydroxide cement

~1.90 Titanium 4.51

Chromium 7.19 Ti6Al4V 4.40

Cobalt chrome alloy

~8.50 Tooth – dentine

2.14

Fluorapatite 3.22 Tooth - enamel

2.97

Glass ionomer cement

~2.10 UHMWPE 0.945

Gold 19.3 Vitreous carbon

1.47

Hydroxyapatite 3.16 Zinc phosphate cement

2.59

Methyl methacrylate

0.94 Zirconia 6.10

Mercury 13.5

TOTAL KNEE REPLACEMENT

1. History

Development of the total knee followed the success of the total hip replacement by Sir John Charnley in the 1960s. He pioneered the use of the polyethylene stainless steel joints fixed to bone with polymethlmethacrylate (PMMA) plastic, often called "bone cement". Today most hip and knee prosthesis are made of cobalt chrome alloy or of titanium. The use of PMMA is fading and many joints are being fixed to bone with new techniques that involve bone ingrowth to the prosthesis. Bone cement still is used widely in the knee with newer techniques to reduce its failure rate. These include mixing under vacuum to prevent air bubbles in the plastic. Bone cement is the same chemistry as Plexiglas (TM) except that bone cement is formed at room temperature and has barium included allowing us to see it on x-rays.

2. Types

Hinge type prosthesis were used initially but had a high rate of failure due to loosening from bone. The unicompartmental knees replace only one part of the joint and have not enjoyed the success of the three compartments or total knee replacement. The modern

devices are minimally constrained. This term means the parts of the knee are not rigidly attached to one another as in a hinge. The successful designs use the ligaments of the knee to hold the knee in place and merely resurface the arthritic joint.

The figure 13 shows total condylar knee prosthesis as it appeared in the 1980s with cobalt-chrome alloy femoral component and high density polyethylene tibial component.

This knee was developed by Install-Burstein and was the standard for total knee replacement for many years.

This knee is still in use with a metal tibial tray, not shown in this photo. The patellar button is also not shown and is round and more than an inch across. The patellar button is made of polyethylene plastic also.

� Figure 13. Total condylar knee prosthesis [19].

3. The Surgery

Total knee replacement is best done in a highly sterile operating room. These are done as the first case in the day because activity in the room stirs up dust. The room is cleaned thoroughly the day before. A clean air filtration system removes airborne dust particles and keeps the air movement horizontal. The surgical team wears sterile gowns that cover the head. These "space suits" protect the patient from debris that could strike the surgeon's face or head and fall back to the wound. The suits also protect the surgeon from contact with bloody material from the bone saw used. Antibiotics are given before surgery to reduce the risk of infection. Total knee replacement requires about 90 minutes of time with the wound open. This means 3 hours in the operating room in


most cases [19].

Figure 14. Surgical picture of a total knee before wound closure [19].

Figure 15. Figure of the Bone Cutting at the knee surgery [19].

ZIRCONIA

Zirconia as a pure oxide does not occur in nature but it is found in baddeleyite and zircon (ZrSiO4) which form the main sources for the material. Of the two of these, zircon is by far the most widespread but it is less pure and

requires a significant amount of processing to yield zirconia. The processing of zirconia involves the separation and removal of undesirable materials and impurities - in the case of zircon - silica, and for baddeleyite, iron and titanium oxides. Typical properties of zirconia are:

- High strength.

- High fracture toughness. - Excellent wear resistance. - High hardness. - Excellent chemical resistance. - High toughness. - Very refractory. - Good oxygen ion conductor. The properties exhibited by zirconia ceramics depend upon the degree and type of stabilisation and on the processing used. Table 11 shows some mechanical properties of zirconia [3].

Table 11. Mechanical properties of zirconia [3].

Property Partially

stabilised Fully Stabilised

Partially stabilised (plasma sprayed)

Density (g.cm-3)

5.7 - 5.75 5.56 - 6.1 5.6-5.7

Hardness -Knoop (GPa)

10-11 10-15

Modulus of Rupture (MPa)

700 245 6-80

Fracture Toughness (MPa.m-1/2)

8 2.8 1.3-3.2

Youngs modulus (GPa)

205 100 -200 48

Poissons ratio 0.23 0.23-0.32 0.25

Thermal expansion (10-6/°K)

8-10.6 13.5 7.6-10.5

Thermal Conductivity (W/m.K)

1.8-2.2 1.7 0.69-2.4

1. Limitations of Zirconia

To date, zirconia’s use has been limited by its loss of strength and its subsequent cracking when subjected to

temperatures of 100-600°C in the presence of water – a process known as hydrothermal degradation. Using state of the art techniques and working at the nanoscale, the research team has inhibited this process by adding trace quantities of materials such as alumina to the zirconia, without compromising its toughness. (One nanometer is one thousand millionth of a metre.) Targeting of the added materials prevents degradation from progressing into the zirconia from its surface [3, 21]. 2. Oxidized Zirconium

Orthopedic surgeons have traditionally delayed joint replacement surgery in patients younger than 65 because they did not expect the materials used to withstand the wear placed on them for longer than 10 to 15 years. Global medical device company Smith & Nephew Inc.'s Orthopaedics Division, in Memphis, TN, has developed


Oxidized Zirconium in response to the medical community's concerns with wear. Smith & Nephew has patented the material for orthopedic use, and the FDA has cleared Oxidized Zirconium for knee implants. Eleven years in development, the Oxidized Zirconium knee is considered an industry-defining technology. Before this technology was developed, nearly 600,000 total knee replacements are performed each year globally. The annual total global knee market is estimated to be $2 billion. Currently, most knee implants are made from a cobalt-chrome alloy that slides against a plastic (polyethylene) bearing. The motion and friction caused by daily living can damage the implant's surface and cause metal and polyethylene wear debris, ultimately causing bone loss and the need for another implant. Because Oxidized Zirconium components are made of a metallic zirconium alloy that is heated to convert the surface to a ceramic (zirconia), the best of both worlds can be achieved. In addition, Oxidized Zirconium contains nondetectable traces of nickel, providing a solution for the more than 20,000 candidates for total knee replacement each year identified as acutely allergic to this metal. Compared to cobalt chrome, Oxidized Zirconium, in wear simulation testing, reduced the rate

of polyethylene wear by 85 percent [10].

3. Zirconium

Orthopedic surgeons at St. Anthony Central Hospital are now pioneering a new knee-replacement prothesis made of zirconium. Through a special process, this metal is heated, then bombarded in an oxygen-enriched environment to yield the second hardest material known, exceeded only by diamonds. The net effect is a femoral component that has the wear characteristics of ceramics without their downside brittleness. Friction is reduced by as much as eight times, which means the replacement joint wears far longer than previous metal models [14].

4. Wear Simulation Comparison of a Zirconia and

a Cobalt Chrome Femoral knee Implant

In recent years the major cause of long-term failure of hip and knee total arthroplasties has been identified as originating with wear particles produced at the interface in the synthetic articulating surfaces. Researchers have tested the hypothesis that a zirconia (zirconium oxide) femur would produce less wear of the counterfacing ultra-high molecular weight polyethylene (UHMWPE) insert than a standard cobalt chrome molybdenum femur of similar design.

The results show a definite reduction in the average steady-state wear rate and the total wear in UHMWPE inserts articulating with the zirconia femurs compared to those articulating with the cobalt-chrome femurs. We speculate that this reduction was due to the increased hardness, scratchresistance and smoothness of the zirconia femurs [20].

JOINT REPLACEMENTS

Total joint replacements of the hip and knee have been performed at St. Anthony Central Hospital since their introduction in the 1970s. During these surgical procedures, mechanical prostheses crafted of specialized metals, ceramics and plastics are used to replace joints irreparably damaged by illnesses (such as rheumatoid arthritis and osteoarthritis) or injury-related conditions (such as vascular necrotitis and post-traumatic arthritis). As new and better designs and materials have become available, outcomes have improved remarkably. Many people having total joint replacement surgery are able to enjoy active, full lives. St. Anthony Central Hospital nurses and therapists conduct a total joint preparatory class twice each month so that those having the surgery know what to expect before, during and after the procedure. Topics covered include pain management, prevention of complications and early mobility [14].

1. Reconstructive Implants

Unique metalworking capabilities and machining techniques, helped create an array of reconstructive implants including large joint replacements, spinal implants, and neurocranial and maxillofacial meshes.

• Hip Systems: First Product that actually is near net shape of a forged hip (Figure 13) and has continued to improve forging and finishing processes for Cobalt Chrome and Titanium hips and acetabular cups to meet the needs for extending implant life.

Figure 13. Forged hip joint [15].

• Knee Systems. The company known as Tecomet's, has such an expertise in forging Titanium, Cobalt Chrome, and Zirconium which provides the solution to the challenges set by complex designs of femoral and tibial components (Figure 14). In June 2001, the company successfully


manufactured the first ever forged Zirconium Femoral Component [15].

Figure 14. Femoral and tibial components [15].

2. Trauma Products

Metal implants remain the dominant form of trauma fixation devices providing superior strength and biocompatibility. Tecomet manufactures a variety of quality trauma products to satisfy the medical industry's need for versatile, cost effective designs.

• Nail Systems. Trauma products for long bones include internal fixation devices (IM Nails). When the orthopedic industry requested strong but lightweight systems, Tecomet responded by developing a proprietary metal forming process to produce hollow titanium fluted nails.

• Plate Systems. From forged plates for long bone fixation to intricate photoetched miniplates for hand surgery, maxillofacial and neurocranial applications, Tecomet's manufacturing and design services support diverse product lines.

• Maxillofacial and Neurocranial Mesh.

Advanced photochemical etching provides the foundation for creating a wide variety of metal reconstruction and fixation implants. Flexible and rigid configurations of fine and coarse meshes allow applications specific to oral and maxillofacial surgery, otolaryngology, neurology, plastic surgery and orthopedics, [15].

3. Capabilities Tecomet excels as a technically strong problem-solving partner prepared to meet the toughest challenges in manufacturing and product development. Through involvement at design inception they have earned a reputation for providing engineered solutions, reducing product launch time and lowering cost. Tecomet's origins are in refractory metals such as molybdenum, tungsten, tantalum and columbium. Extensive experience in forging, machining and the development of technologies to process high strength

materials such as titanium, zirconium, kovar, nickel and cobalt super alloys enables us to create products with demanding performance requirements, [15].

4. Technologies and Procedures

The primary concern surgeons seem to have regarding metal-on-metal implant procedures is elevated serum chromium levels. “To date, there have been no reports of cobalt toxicity in patients with elevated levels of serum cobalt in association with metal-on-metal total hip replacements,” says Josh Jacobs, MD, of Rush Memorial University of Chicago Hospitals. However, Jacobs adds that the literature is incomplete and the necessary studies have not been conducted to determine whether these elevated levels are a long-term concern. The advantage of the metal-on-metal implant is its longevity, which minimizes the need for a later revision.

The ceramic-on-ceramic implants have enjoyed popularity outside the United States for years; however, they are associated with higher incidents of fracture. James T. Caillouette, MD, attending surgeon at Hoag Memorial Hospital, Newport Beach, Calif, and an assistant clinical professor at the University of California at Irvine, observes, “There was a nexus of unfortunate events surrounding the FDA [approval] proceedings. They were trying to protect us. Under the circumstances, the incidents that fostered concern at the FDA were comparatively minute in terms of the number of procedures that occur without complications.”

Metal-on-cross-linked polyethylene shows promise, according to orthopedic surgeon Richard A. Berger, MD, of Rush Memorial University of Chicago Hospitals. But there is some evidence in the presence of abrasive particles that wear may, in fact, be accelerated with cross-linked polyethylene. Some studies have indicated that the wear debris caused by repeated contact between the articulating surfaces has been a lingering clinical concern. “The most cost-effective procedure is cross-linked polyethylene on a chrome cobalt head,” says Berger. “It's [the procedure] most [physicians and patients] are choosing across the country.”

The latest FDA-approved innovation in hip arthroplasty is the oxidized zirconium implant, a new material that combines the advantages of scratch resistance and extremely low wear but without the potential for fracture or high metal ion levels. Oxidized zirconium femoral heads are made from zirconium metal that has been oxidized such that the surface of the implant is a ceramic zirconia. The oxidized surface has proven to be extremely hard and abrasion resistant. These laboratory tests have validated that oxidized zirconium resists abrasion, minimizes wear, and has no demonstrable risk of breakage or delamination, [2].


5. Biologically Compatible

Suboptimal alignment of hip prostheses—for example, excessive vertical positioning of the acetabular component—increases wear, especially near the periphery of the component. Anatomic restoration of the hip center of rotation and offset and avoidance of impingement are associated with decreased wear. Optimal surgical technique involves stable fixation to minimize interfacial motion and avoidance of residual particles that could potentially contribute to third body wear. But not all new techniques have built on the materials and methods first pioneered in the 1960s. Recent studies of positive outcomes of hip procedures using noncemented tapered stems have made them among the most favored of orthopedic physicians. Fresno, Calif-based orthopedist D. Kevin Lester, MD, who specializes in minimally invasive hip procedures, and is an assistant clinical professor at the University of California, San Francisco, is an enthusiastic proponent of the cementless tapered titanium femoral prosthesis. “The cementless, collarless hip implant is totally compatible with a minimally invasive procedure, and has zero failures due to loosening, and 12% chance of improvement,” he says. With improved materials and techniques have come less invasive procedures, which have aided in the success of hip arthroplasty,[2].

6. Orthopedic Biomaterials

Producing a material that can function intimately with living tissue, with minimal adverse reaction, is quite a challenge for engineers and scientists. Biomaterials are designed to perform specific functions in the body and, at times, are used to replace parts of living systems. Some common implants include knee and hip joint replacements, spinal implants, and bone reinforcement devices. Also popular are artificial heart valves, soft tissue replacements, and a variety of dental implants. Each of these devices must be constructed of special materials that are uniquely suited for their respective tasks. Properties such as mechanical integrity, corrosion resistance, and biocompatibility must be evaluated for any biomaterials, [8].

7. Materials for knee replacements

Unlike hip replacement devices there is currently little choice in materials for knee replacements. A three year European Community funded a program to explore the use of ceramic materials in knee arthroplasty has demonstrated that ceramic on polyethylene combinations reduce polyethylene wear compared to existing metal on polyethylene bearings. Product features include:

• Metal backed ceramic femoral component.

• Zyranox ® zirconia on polyethylene bearing.

• Existing, clinically proven fixation methods.

• Patented braze technique for joining ceramic to metal.

• Ceramic resurfacing available for any current knee system [12].

8. Characteristics of Materials Used in Orthopedics

a. Fracture fixation :

i. Stainless steel

• Iron based alloy containing chromium, nickel, molybdenum. Usually annealed, cold worked or cold forged for increased strength. A range of strength and ductilities can be produced.

• Strong.

• Cheap.

• Relatively ductile therefore easy to alter shape. Useful in contouring of plates and wires during operative procedures.

• Relatively biocompatible.

• The chromium forms an oxide layer when dipped in nitric acid to reduce corrosion and the molybdenum increases this protection when compared to other steels.

• Can still undergo corrosion if carbon gets to the surface.

• High Young’s modulus - 200 GPa (10× that of bone) so can lead to stress shielding of surrounding bone which can cause bone resorption.

• Used in plates, screws, external fixators, I.M. nails.

• Composition of 316L Stainless Steel: Iron- 60%, Chromium- 20% (major corrosion protection), Nickel- 14% (corrosion resistance), Molybdenum- 3% (protects against pitting corrosion), Carbon- 0.03% (incr. strength), Manganese, Silicon,P,S,- 3% (control manufacturing problems).

ii. Titanium and its alloys

• Excellent resistance to corrosion .

• Young’s modulus approximately half that of stainless steel, therefore less risk of stress protection of bone, stress riser at end of plate or nail.

• More expensive than stainless steels.

• Poorer wear characteristics than others, therefore not considered suitable as a load bearing surface these days.

• Can be brittle i.e. less ductile than stainless steel, but more ductile titanium alloys being produced.

• Can be as strong as stainless steels.

• Used in plates, screws, I.M. nails, external fixators. Useful in halos as more MR scan compatible than other metals.


iii. Adhesives

• Not common in orthopaedics but potentially useful in small fragment fixation, controversial.

• Need to 1. Have sufficient bond strength 2. Be able to bond to moist surfaces 3. Permit healing across the bond line 4. Be sterilisable.

• Bone cement does not count as an adhesive.

• Cyanoacrylate-Experiments on rabbits produced poor results, long term fate not addressed.

• Fibrin- Conclusions of experiments are that fibrin adhesives are only suitable for fracture fragments with considerable inherent stability or being non load bearing.

iv. Biodegradable polymers

• Potential advantages o Hardware removal not necessary,

reducing morbidity and cost. o Stiffness of polymer decreases as

stiffness of fracture callus increases.

o Can possibly be used in future for controlled release of antibiotics or stimulants to healing .

• Requirements o Adequate mechanical strength for

the application o sufficient strength over a sufficient

period of time to maintain enough stability for the fracture to heal and prevent loss of reduction

o Degradability into products that are mot harmful.

• Examples o Polyglycolic acid o Polylactic acid o Copolymers

• Only about 1/20th the stiffness and strength of stainless steel

• Used in ankle fractures with poor results

• Used in phalangeal fractures with better results

b. Materials Used in Joint Replacement Surgery

i. Stainless Steel

• Now rarely used in new designs

• Because Young’s modulus high, need to be inserted with a lower modulus polymer cement for fixation, to prevent stress shielding of the surrounding bone.

ii. Cobalt Chrome

• 30-60% cobalt, 20-30% chromium,7-10% molybdenum + nickel.

• Stronger and more corrosion resistant than stainless steel.

• Young’s modulus higher than stainless steel (250 cf 200 GPa). Stress shielding a theoretical risk. Usually fixed with cement.

iii. Titanium Alloys

• Most common combination is Ti6Al4V

• Strong and corrosion resistant

• Young’s modulus 110GPa (less than cobalt chrome & stainless steel), therefore often used for cementless joint replacements.

• Poorer wear characteristics.

• Ultimate Strength: Stainless Steel > Titanium; Yield Strength (permanent deformation): Titanium > Stainless Steel

• Ti13Zr13Nb is stronger and has lower Young’s modulus.

• Theoretically, may favour bone apposition and bone ingrowth more than cobalt chrome, but no difference found clinically.

iv. Polyethylene

• UHMWPE- Ultra high molecular weight polyethylene. A polymer of ethylene.

• Molecular weight 2-6 million.

• 90% success rates at 15 years with metal on polyethylene (therefore the gold standard).

• The weak link of any total joint replacement.

• Osteolysis produced due to polyethylene wear debris causes aseptic loosening.

• Submicron particles found in periprosthetic tissues when polyethylene wear present.

• Factors affecting polyethylene wear: o Material polymorphism - Ziegker

process used to produce polymerisation. Consolidation produced by Ram extrusion or compression moulded. Component machined from these blocks. Bankston et al - linear wear rate 0.05mm /year for compression moulded, 0.11mm / year for ram extruded.

o Gamma sterilisation in air produces chain scission by oxidation. Companies now vacuum pack and sterilise the implants.

o Thin polyethylene- increases wear, due to increased fatigue wear if thickness less than 6-8mm.

o Polyethylene should not directly touch bone in hip replacement- increases wear.


o Conformity- increased conformity reduces stresses ( particularly relevant in TKR

o Materials used- Titanium bearing surface increases wear, ceramics reduce wear.

o Size of femoral head:

� Large femoral head, causes increased sliding distance at joint surface and so increased adhesive wear. Therefore increased volumetric wear (Theory behind Charnley's LFA 22mm head).

� Small femoral head causes increased stresses (increased fatigue wear or penetrative wear) Trade off ideal femoral head size is 28mm.

o Reduced offset of femoral prosthesis causes increased fatigue wear, as joint reaction force higher.

o Three body wear:

� Due to cement particles, metallic particles.

� Can be reduced by careful surgical technique.

� Can be increased with modularity of implants.

v. Ceramics

• Strong ionic bonds between the metallic and nonmetallic components.

• Very strong.

• Very stiff.

• Very biocompatible.

• Very hard, therefore good wear characteristics.

• But very brittle.

• Also difficult to process due to very high melting points therefore expensive.

• Bioinert e.g. Alumina, Zirconia, used for surface replacement.

• Bioactive e.g. hydroxyapatite and glass- used for coating joint replacements for osseo integration between bone and implant.

vi. Hydroxyapatite coating of THR

• Ca10 (PO4)(OH)2 coated onto metal surface, usually onto a porous surface

• Usually 50-150µm thick. (Too thin can be resorbed, too thick can flake off during insertion of implant).

• Is thought to be osteoconductive

• Not known how long it takes to resorb and how stable the implant is after resorption

• Good results at 5 years (99% survival) according to Norwegian Arthroplasty register

• Some worry about increased three body wear on polyethylene

• The particles of HA may also stimulate osteolysis.

vii. Bone cement

• Polymethylmethacrylate introduced 30 years ago

• No other fixation principle has given better long term results

• Polymerised methyl methacrylates

• mixed from powder polymer and liquid monomer in theatre usually in a vacuum to reduce porosity and increase strength. Powder also contains catalyst ( benzoyl peroxide).

• Xray contrast medium (barium sulphate)

• Colour (chlorophyll)

• Stronger in compression than tension, weakest in shear

• Exothermic reaction producing heat, can lead to bone necrosis

• Leakage of monomer during polymerisation may induce endothelial damage locally leading to thrombus formation and distal hypotension through effects in pulmonary vascular bed

• Cementing techniques have changed from finger packing to retrograde filling with a cement gun under maximal pressurisation - has not been evaluated in randomised clinical studies

• Controversy about type of bond between implant and cement- should it be maximal with roughening of implant surface or minimal with a polished surface [11].

COMPARISON OF PROPERTIES BETWEEN

STAINLESS STEEL 316L WITH OTHER

BIOMATERIALS

Table 12 shows a comparison of properties between stainless steel 316 L with the materials studied in this research


Table 12. Properties of biomaterials [22, 23].

Material Modulus of elasticity ( GPa )

Shear modulus of elasticity ( GPa )

Poisson’s ratio

Yield stress Mpa

Stainless steel

200 82 0.27-0.30 min- 170

Cobalt chrome

230 .30 413

Zirconium 200 70 0.22 2000

Titanium 100-120 39-45 0.33 760-1000

Calcium Sulfate

18-21 6-10 .305 -------

Calcium Phosphate

18-21 7-13 .315 -------

SUMMARY

The most common materials used in orthopedics are: titanium, zirconium, cobalt-chrome, calcium phosphate, and calcium sulfate and stainless steel 316-L. Titanium is used primarily for the loading faces which include the pin structure, fabrication of plates and femoral stems. The Modulus of Elasticity of Titanium is much lower than Stainless Steel 316-L, having a numerical difference in value which ranges from 80-100 GPa. The Shear Modulus of Elasticity is also lower than the particular value of Stainless Steel 316-L, this difference is about 37 GPa. The difference in Poisson’s Ratio is just about 3 decimal units, but the yield stress of Titanium is much more higher in comparison to that of Stainless Steel 316-L. The difference expressed is more than 600 MPa for the yield stress. In conclusion Stainless Steel 316-L is much stronger, but that is not always good because stress rises at end of plate or nail. Titanium possesses a lower ultimate strength than Stainless Steel 316-L but its yield strength is much more, this is what causes permanent deformation of the material, and Stainless Steel is easily expected to recover its normal state than Titanium and it’s Alloys.

Cobalt-Chrome which is a cobalt alloy has a Modulus of Elasticity 230 GPa, when compared to the Modulus of elasticity of Stainless Steel gives us a difference of 30 GPa. In this particular case cobalt chrome’s Modulus of elasticity is higher than Stainless Steel. The Poisson’s Ratio of both are very similar, they both are near 0.30. The yield stress that cobalt can support is 413 Mpa. When compared to Stainless Steel 316-L the difference obtained is near to 243 Mpa. In conclusion, this material seems to be better than Stainless Steel 316-L, but the only disadvantage is the price and the facility to find it. Zirconium’s Modulus of Elasticity is 230 GPa which is very close to the Modulus of Elasticity of Stainless Steel. The Shear Modulus of Elasticity is 70 GPa, when compared to Stainless Steel 316-L it gives a difference of 12 GPa. The Poisson’s Ration of zirconium is 0.33 and the difference between both is

only from 3 to 4 units. The materials seem similar but the yield stress of both materials is different. Zirconium has a 2,000 Mpa yield stress value and stainless steel has only 170 MPa. This particular property gives the material is the maximum stress it could hold and return to its original state. In the Bone Grafting face, the predominant materials are calcium sulfate and calcium phosphate. The Modulus of Elasticity of calcium phosphate is in the range from 18-21 MPa. This material compared to stainless steel 316-L has a differential value of more that 180 MPa which clearly states that is not as elastic as stainless steel, which by definition makes it much more brittle. The Shear Modulus of Elasticity of Calcium Phosphate is about 7-13 GPa, and the differential value is more than 68 GPa. The Poisson’s Ratio fluctuates in the same range. The yield stress, which is the ability of the material to recover or to return to its normal state, has a value so low that is not taken into consideration. The yield stress of the Stainless Steel 316-L is much more dominant. Calcium sulfate has very similar properties as calcium phosphate, which also makes it very brittle. The Modulus of Elasticity of calcium sulfate ranges from 18-21GPa which compared to Stainless Steel 316-L the difference in values is very close to that of calcium phosphate which is in the range from 170-180 MPa. By definition calcium sulfate is a very brittle material. The difference in numerical values of the Shear Modulus of Elasticity also fluctuates from 68 GPa and over. The Poisson’s Ratio is almost in the same range, and has a numerical value of 0.315. Since the material is so brittle the yield stress is not taken into much consideration, as in the comparison with calcium phosphate, stainless Steel is much more dominant in this precise property. As observed in bone grafting the materials used are brittle but very strong and stiff, because of this they are used in joint replacement and osseo integration between bone and implant, some sort of a refilling process, and not as Stainless Steel which is used in other orthopedic surgery processes.

ACKNOWLEDGEMENTS

Our thanks to Dr. Megh R. Goyal for his guidance.

REFERENCES

1. http://www.orthobluejournal.com. Szpalski, Marek and Gunzburg, Robert. Applications of Calcium Phosphate-Based Cancellous Bone Void Fillers in Trauma Surgery.

2. http://www.orthopedictechreview.com/issues/mayjune03/pg36.htm .

Gordon II, MA, W.A. A Technical Evolution.


3. http://www.azom.com. Azom-Metals, Ceramics ,Polymers, Composites An Engineer Resource. 2004. A to Z of Material.

4. www.azom.com Biomaterials: An Overview.

5. http://www.suri.org/3pubs/ttoday/fall95/implant.htm

Blanchard, Cheryl R. Biomaterials: Body Parts of the Future. 6. http://www.orthobluejournal.com

Parikh, Shital N. Bone Graft Substitutes in Modern Orthopedics.

7. http://www.karger.com/gazette/65/lidgren/art_5_0.htm

Lindgren, Lars. Bone Substitutes. 8. http://www.cheresources.com/newparts.sht

ml. Building New Body Parts. 9. http://www.biomed.tamu.edu 10. http://www.goodsamdayton.com/knee.htm

Good Samaritan Hospital First in Area to Introduce New Technology for Knee Implants.

11. http://www.orthoteers.co.uk/Nrujp~ij33lm/orthomat. htm

Implants & Materials in Orthopedics. 12. http://www.bioceramics.com/knee1.htm

Knee Replacements 13. Hasting,G.W and D.F. Williams.,1980.

Mechanical Properties of Biomaterials. John Wiley & Sons.

14. http://www.stanthon4central.org Total Joint Replacement.

15. http://www.orthosupplier.com/players/4b.htm Tecomet

16. http://www.worldortho.com/history.html Brakoulias, Vlasios. The History of Orthopedics.

17. 17.http://www.ae.msstate.edu/vlsm/materials/alloys/titanium.htm. Titanium Alloys and Their Classification.

18. www.azom.com Titanium and Titanium Alloys as Biomaterials.

19. http://www.ucbones.com/total_knee_replacement.htm

Total Knee Replacement 20. http://www.jbjs.org/ORS_2001/pdfs/1101

Wear Simulation Comparison of a Zirconia and a Cobalt Chrome Femoral Knee Implants.

21. www.azom.com Zirconia (ZrO2) – Is Zirconia a Viable Alternative to Steel.

22. http://www.efunda.com 23. www.matweb.com 24. http://www.btec.cmu.edu/reFramed/tutorial/

mainLayoutTutorial.html The Need for Bone Substitutes

GLOSSARY

Alignment - Positioning the femur and tibia so as to allow proper articulation at the knee joint. Allograft - A graft (occasionally bone) taken from a human being and implanted in another. Alloy - A mixture of two or more metals. Anatomic - Relating to the structure of an organism. Often used to describe a prosthesis, which closely resembles or duplicates the shape and size of a normal part of the body. Anatomic Axis - The axis formed by an imaginary line down the center of the femoral canal. Usually 5-7 degrees off the mechanical axis. Ankylosis - The fusion of a joint. Arthrodesis - Surgical fixation of a joint. Articulating Surface - Implant or bone surfaces which touch each other. Typically used in referring to the polyethylene tibial surface or patellar surface. Autograft - A graft (sometimes bone) taken from a patient and reimplanted in another part of his/her own body. AVN - Avascular necrosis (particularly death of bone through lack of blood supply).

Biocompatibility - Referring to the degree of tissue or systemic reaction caused by a foreign material in the body. Biomechanics - Relating to the forces that act on the joint, and their effect on the joint. Bone Ingrowth - The process of bone growing into the pores of a porous implant for enhanced fixation. Bone resorption - A remodeling of bone due to a lack of stress through an applied load. A common result of stress shielding, where bone located in an area that is shielded from stress is absorbed by the surrounding bone that is under stress. Also called Stress Relief Osteoporosis.

Calcaneous - The heel bone Calcar - An area in the media region of the proximal femur which is characterized by very dense cortical bone. Caliper - An instrument used to determine thickness, diameter, or width. Cancellous Bone - Spongy bone composed of a loose latticework of bony traveculae and bone marrow within the inerspace found in the enlarged ends of long bone. Cannulated - An open-ended passageway through which a wire or pin may be passed. Cast - The giving of shape to metal by pouring it in liquid form into a mold and allowing it to solidify. Closed Procedure - Done with the use of image intensification but without the need of an incision at the fracture site. Collateral Ligaments - The strong stabilizing ligaments located on both the medial and lateral sides of the knee. Comminuted Fracture - One in which there are several definite disruptions in the bone, creating two or more fragments Compartment - A combination of two surfaces which articulate with each other. Compression - Adjustment of an external fixator to provide closer bone-to-bone contact at a fracture site; or the application of centrally directed forces (forces applied towards the middle of the instrumented areas. Contracture - Abnormal shortening of muscle tissue,


rendering the muscles highly resistant to stretching. Cortex - The outer surface of a bone or organ. Countersink - Instrument used to form a flaring depression around the top of a drilled hole. Insertion of an implant beneath the cortical surface of the bone. Cruciate ligaments - Two strong stabilizing ligaments which cross between the condyles of the knee. The anterior cruciate ligament runs from the back of the femur to the front of the tibia. The posterior cruciate ligament runs from the front of the femur to the back of the tibia.

Deformity - A change in shape or form. Delayed union - An abnormal lag in the reunion of fractured parts of a bone. Diaphysis - The shaft of a long bone, as distinguished from the extremities and outgrowths. Dislocation - Displacement of two bones from their normal location at a joint. Distal - Situated or directed farther away from the center of the body or some part of the body. Distraction - A type of dislocation in which the bones of a joint have been separated without damage to their ligament. Endoprosthesis - Repair of the femoral head only where the acetabulum is left intact. Endosteum - Membrane lining bone in the medullary cavity. Epiphyseal Growth Plates - The ends of long bones at which growth in length mainly takes place. Epiphysis - Pertaining to the end of long bones, usually wider than the shaft. Eversion - The act of turning outward; opposite of inversion. Exchange Nailing - A procedure where an existing IM nail is removed and replaced by a new nail. Extension - The act of extending or straightening a limb.

FDA - Food and Drug Administration. A federal agency which regulates the sale and distribution of surgical implants. Fatigue Strength - The ability of a material to withstand cyclic stress. Femoral Head - The "ball" portion of the hip joint, located at the proximal end of the femur. Femur - The long bone between the hip and the knee. Fibula - The long thin outer bone of the lower leg. Fixation - The process of making an object hold firm or fast. Flexion - The act of flexing or bending a joint so that the bones forming it are brought toward each other. Fracture Callus - An unorganized meshwork of woven bone developed on the pattern of the original fibrin clot, which is formed following fracture of a bone. Frontal Plane - A plane which divides the body into dorsal and ventral parts. Also called the coronal plane. Fusion - The joining together of two structures (such as the joint space between the femur and the tibia).

Humerus - The upper arm bone.

Image Intensification - Fluoroscopic radiographic monitoring of fracture site and procedure while in progress. The x-ray machine is commonly referred to as a C-Arm because of the shape of the machine. Infrapatellar ligament - A ligament located directly below the patella. Intercondylar notch - An indentation between the two condyles of the distal femur. Interfragmental compression - To apply compression to two or more bone fragments using bone screws. Interlocking nail - An intramedullary nail that is designed to accept cross-screws both proximally and distally to improve rigidity and rotational stability in early fracture healing. Screws can be phased out to improve load sharing. Intracondylar notch - The recess between the femoral condyles of the knee joint which housed the two cruciate ligaments. Intradedullary - Within the medullary, or narrow, cancellous canal region of a long bone. Intramedullary Canal - The marrow cavity of a bone. Intraoperative - Performed during surgery. Ligament - A tough band of white fibrous connective tissue that links two bones together at a joint.

Malleable - Capable of being extended or shaped by pressure. Malunion - Growth of fractured bone fragments in a faulty position, forming an imperfect union. Marrow - Soft central part of a bone. Mechanical axis - The axis formed by a line which passes through the center of the hip, the center of the knee, and the center of the ankle. Medial - Situated or directed toward the midline of the body or one of its parts. Medial malleolus - A protuberance on the lower end of the tibia at the ankle. Medullary canal - The inner portion of a bone which is filled with bone marrow. Modularity - Referring to an implant system where a specific component is made up of two or more detachable parts. Modulus of elasticity - A measure of the ability of material to return to its original shape without any permanent deformation when stress is applied. Moment - The tendency of a force to produce a rotation at a certain point. Nonunion - Failure of fractured bone to heal completely. Oblique - Slanting; between horizontal and vertical in direction. Osteochondritis dissecans - A degenerative process to the articular cartilage by which a separation of the articular cartilage from the sub-chondral bone occurs.

Patella (kneecap) - A sesamoid bone (a bone contained within a tendon) that essentially acts as a "pulley


mechanism" to complete patella femoral joint. Pelvis - A basin or basin-like structure. Percutaneous - Effected through the skin. Periosteum - A thick fibrous membrane covering the surface of bone except at points of articulation. Posterior cruciate ligament - An internal ligament of the knee joint originating from the medial femoral condyle and inserting into the notch of the posterior tibia. Function: stabilizes the tibia on the femur Primary - A procedure where an implant is implanted where there has not previously been an implant. Pronate - To turn palm down. Prone - Lying with face downward. Prosthesis - An artificial substitute for a missing part of the body. Radiolucent - Permitting the passage of x-rays. Radiopaque - Not permitting the passage of x-rays. Radius - A bone in the forearm. Range of motion - The arc created by the flexion of a limb at the joint, usually expressed by degrees. Reamer - An instrument used to enlarge or shape a drilled hole; such as the intramedullary canal of a long bone. Reduction - The restoration of a dislocated part of the body to its normal position. Reproducible - A procedure or instrument system that is designed to achieve the same basic results from all surgeons and in all cases. Resection - The surgical removal of part of a bone. Retractor - An instrument used to expose tissue or bone. Retroversion - A backward turning; opposite of anteversion. Rongeur - Biting forceps for cutting bone. Rotation - Turning about an axis or a center.

Sawbones - A plastic replica of a bone which is used for testing, practicing or demonstrating a surgical procedure Segmental Fracture - Divides a long bone into more than two roughly cylindrical segments (also a specific type of comminuted fracture). Soft Tissue - Any of the tissues other than bone that surround or are within the joint, including muscles, tendons, ligaments, etc. Stability - The degree of resistance to forces tending to cause motion or change of motion. Stress - A force or action placed on a surface or material. Stress riser - A point or area where stress is concentrated due to an uneven distribution of stress. If an implant fails, it fails at a stress riser. Subluxation - A partial dislocation of a joint, so that the bone ends are misaligned but still in contact.

Tap -Instrument used to cut threads into the bone for a bone screw. Tendon -A tough band of white fibrous connective tissue that links a muscle to a bone. Tensile strength -A measure of resistance to tensile stress.

Tensile stress -Stress generated by a force which tends to elongate or stretch a body or structure. Tibia -The inner and larger bone of the lower leg. Tibial plateau -The upper end of the tibia capped with articular cartilage, articulating with the femur, making up the lower half of the knee joint. Tibial spine -A projection on the top surface of some tibial components between the two halves of the tibial plateau. Tibiofemoral -Pertaining to both the tibia and the femur. Titanium -A metallic element used to make surgical implants. Titanium is very biocompatible and offers a lower modulus of elasticity than many other metals. Torque -A force that produces or tends to produce rotation or torsion. It is a measurement of an instrument's capacity to do work or to continue to rotate under resistance to rotation. It is expressed in inch-ounces or inch-pounds. Torsion -Act of twisting or condition of being twisted. Torsional stress -Stress generated by a force which tends to rotate or twist a body or structure. Trocar -A sharp obturator used to assist the insertion of the bushings through soft tissue. Trochanter -Either of the two processes below the neck of the femur. The greater trochanter is a broad, flat process at the upper and lateral surface of the femur to which several muscles are attached. The lesser trochanter is a short conical process projecting medially from the base of the neck of the femur. Tuberosity -A large rounded eminence on a bone.

Valgus - Deformity in bone, such as when the knees are close together (knock-knee), with the ankle space increased. Bent outward. Varus - Deformity in bone, such as when knees are bowed out, and the ankles are close in. Turned inward.


Appendix I: NUMERICAL EXERCISES

Exercises:

Chapter 1, Problem 1.2-4 A circular aluminum tube of length L=400mm is loaded in compression by forces P (see figure 1). The outside and inside diameters are 60mm and 50 mm, repectvely. A strain gage is placed on the outside of the bar to measure normal strains in the longitudinal direction. (a) If the measured strain is ε=540 x 10 -6, what is the shortening δ of the bar? (b) If the compressive stress in the bar is intended to be 40MPa, what should be the load P? Discussion For the application in the biomechanics in orthopedics, the geometry of the tube could be considered circular (ideal behavior) and the material is titanium. The material changed because the bones of the body are not made of aluminium, but as we learn, a metal as titanium could be implanted. The dimensions as length and diameters could be the same, because the anatomy of the body allows these dimensions. The forced of compression would be the compression of another two bones in connection, with the weight of the body within. According to the changes made the question (a) could be rephrase with different value and the part (b) could stay the same.

Figure 1

( ) (( )

KNP

MPaP

MPaP

P

MPaIf

mmoL

oLL

EAPL

mmmoutd

mmmind

mmmL

6.34

0346.0

205.0

206.04/40

?

40,

220.04

1020.2

06.060

05.050

4.0400

610550

=

=

−=

=

=

=−

×==

=

=

==

==

==

−×=

π

σ

εδ

δ

δ

ε

Chapter 2, problem 2.4-17 A trimetallic bar is uniformly compressed by an axial force P=2000 lb applied through a rigid end late (see figure 2). The bar consists of a circular steel core surrounded by brass and copper tubes. The steel core has a diameter 0.4in., the brass tube has outer diameter 0.6in., responding moduli of elasticity are Es=15 x 106psi, and Ec=18 x 106 psi. Calculate the compressive stress σs, σb, and σc in the steel, brass, and copper respectively, due to the force P.

Discussion For the application of this problem in the biomechanics in orthopedics, the geometry could be considered circular, as stated for ideal behavior. The material would be changed for the zirconium alloy that is also used in the orthopedics applications. The load could be the weight of the body and the base the floor where the load is applied. The load could be of P=150 lb. The diameters could be changed for diameters less than 3 in, assumed as 2 in. The Ez=2.9 x 107 psi is the modulus of elasticity of zirconium, and having a length of 10 in. According to the new values and new changes the question should be arranged.

P = 150 lb d = 2 inch E = 2.9 * 107 Psi ∑ F = p σ = P/A = 47.75 Pa δ = PL/EA = (150)(10)/( 2.9 * 107)(π12) = 1.65 *10-5 inch

Chapter 3, problem 3.4-1 A stepped shaft ABC consisting of two solid circular segments is subjected to torques T1 and T2 acting in opposite directions, as shown in the figure. The larger segment of the shaft has a diameter d1=2.5 in. and length L1= 25 in.; the smaller segment has diameter d2= 2.0 in. and length L2= 18 in. The material is steel with shear modulus G=11 x 106 psi. If the torques is T1=9000lb-in. and T2=4000lb-in., calculate the following quantities: (a) the maximum shear stress τmax in the shaft, and (b) the angle of twist φC (in degrees) at end C.


Discussion For the application in the biomechanics in orthopedics, the geometry of the tube could be considered circular as ideal behavior. The differences in diameters could be seen in the arms bones which vary in dimensions and are fixed to the shoulder. The length of the shaft should be change in order to be more truthful with the anatomy of the body. The torques applied in the shaft could be seen as two different torques applied to the different parts of the arm at the same time. The material is titanium with modulus of elasticity known (see table 3). According to the new values and new changes the question should be arranged.

radBCABC

psiBC

radBC

psi

ininlb

PGITL

BC

radAB

psi

ininlb

PGITL

AB

psid

TBCT

inlbTBCT

psid

TABI

inlbTTAB

T

inL

inL

ind

ind

psiG

inlbT

inlbT

610935.1

55.9max

71068.8

59.759,278,1715

)24)(32/)(61011(

)10)(15(

610067.1

67.471,184,4245

)54.2)(32/)(61011(

)9)(5(

55.913.25

240

3)2(

)15(163

16

152

63.109.49

80

3)5.2(

)5(163

16

5101521

102

0.91

0.22

5.21

61011

152

101

−×=+=

==

−×==

×

−==

−×==

×

−==

====

−==

====

−=−=−=

=

=

=

=

×=

−=

−=

φφφ

ττ

φ

πφ

φ

πφ

ππ

ππ

Chapter 4, problem 4.3-15 Beam ABCD represents a reinforced-concrete foundation beam that supports a uniform load of intensity q1=2100 lb/ft (see figure). Assume that soil pressure on the underside of the beam is uniformly distributed with intensity q2. (a) Find the shear force VB and bending moment MB at point B. Discussion

For the application in the biomechanics in orthopedics, the geometry of the tube is not taken into consideration for the analysis. The load supported would be suggested as the weight supported by a foot. The load applied and uniformly distributed would be q1 =150 lb/ft and the length of the foot will be 1 ft. The load applied as q2 could be considered as the uniformly load exerted by the floor when the weight load is applied. According to the new values and new changes the question should not be arranged.

q2(7) = q1(1) q2 = 21.43 lb/ft

Vb, Mb in b Vb = (21.43) (3) = 21.43lb a)Mb = (21.43)(3)(1.51) = 96.44 lb-ft

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selected biomaterials

biomaterials scientist

examples of biomaterials