Сraig. Dental Materials

products on the local and systemic tissues; and

(3) the interfacial zone between the implant and tissue. Regarding the ultrastructure of the implant/tissue interface, it is important to understand that, although this zone is relatively thin (on the order of ~ngstroms),the constituents of the zone-heterogeneous metallic oxide, proteinaceous layer, and connective tissue-have a substantial effect on the maintenance of interfacial integrity. Furthermore, interfacial integrity depends on material, mechanical, chemical, surface, biological, and local environmental factors, all of which change as functions of time in vivo. Thus implant success is a function of biomaterial and biomechanical factors, as well as of surgical techniques, tissue healing, and a patient's overall medical and dental status.

The physical and material factors affecting implants may be summarized as follows:

(1) materials and material processing; ( 2 ) mechanisms of implant/tissue attachment; (3) mechanical properties; ( 4 ) implant design; (5) loading type; (6) tissue properties; (7)stress distribution;

(8)initial stability and mechanisms of enhancing osseointegration; (9) biocompatibility of the implant materials; and (10) surface chemistry, mechanics, and bone-bonding ability of the implant.

Implant Materials and Processing In general, two basic classes of materials-metals and ceramics-are used in implantology, either alone or in hybrid fashion. Metallic implant materials are largely titanium based-either c.p. Ti or Ti-6A1-4V. However, it is essential to note that the synergistic relationships between processing, composition, structure, and properties of the bulk metals and their surface oxides effectively leaves more than two metals. Casting, forging, and machining to form near-net shaped end products can alter the bulk microstructure, surface chemistry, and properties. Similarly, densification of ceramics and deposition of ceramic and metal coatings by hot isostatic pressing or sintering can change bulk and surface composition, structure, and properties. Thus the many material processing sequences necessary to yield the end-stage dental implant have a strong influ-

ence on the properties and functionality of the implant, primarily through temperature and pressure effects.

Although cobalt-based alloys have been used experimentally in dentistry, clinically used metallic dental implants are almost exclusively titanium based. The attributes of titanium, namely, corrosion resistance and high strength, have been discussed previously.

The initial rationale for using ceramics in dentistry was based on the relative biological inertness of ceramics as compared with that of metals. Ceramics are fully oxidized materials and therefore chemically stable. Thus ceramics are less likely to elicit an adverse biological response than metals, which oxidize only at their surface. Recently, a greater emphasis has been placed on bioactive and bioresorbable ceramics, materials that not only elicit normal tissue formation, but that may also form an intimate bond with bone tissue and even be replaced by tissue over time.

Mechanisms of Implant/Tissue Attachment Various surface configurations have been proposed as means of improving the cohesiveness of the implant/tissue interface, maximizing load transfer, minimizing relative motion between the implant and tissue, minimizing fibrous integration and loosening, and lengthening the service-life of the construct. The concept of osseointegration around screw-threaded implants represents a situation of bone ongrowth. An alternative method of implant fixation is based on bone tissue ingrowth into roughened or threedimensional porous surface layers. Such a composite system has been shown to have a higher bone/metal shear strength than have other types of fixation. Increased interfacial shear strength results in a better stress transfer from the implant to the surrounding bone, a more uniform stress distribution between the implant and bone, and lower stresses in the implant. In principle, the result of a stronger interfacial bond is a decreased propensity for implant loosening. A progression of surfaces from the lowest implant/tissue shear strength to the highest is as follows: smooth,

textured, screw threaded, plasma sprayed, and porous coated.

Enhancing Osseointegration Because of the necessity of developing a stable interface before loading, effort has been placed on developing materials and methods of accelerating tissue apposition to dental implant surfaces. Materials developments that have been implemented into clinical practice include using surfaceroughened implants and ceramic coatings. Other, more-experimental techniques include electrical stimulation, bone grafting, and the use of growth factors.

Important examples of bioactive materials are bioactive glasses and glass ceramics and calciumphosphate ceramics. Bioactive glasses and glass ceramics include Bioglass, which is a synthesis of several glasses containing mixtures of silica, phosphate, calcia, and soda; Ceravital, which has a different alkali oxide concentration from that of Bioglass; and a glass ceramiccontaining crystalline oxyapatite and fluorapatite (Ca,,[P04]6[0,F,]) and P-wollastonite (SO,- CaO) in a Mg0-Ca0-Si0, glassy matrix (denoted glass-ceramic A-W). Calcium-phosphate ceramics are ceramic materials with varying calcium/ phosphate ratios, depending on processinginduced physical and chemical changes. Among them, the apatite ceramics, one of which is hydroxyapatite (HA), have been studied most.

The impetus for using synthetic HA as a biomaterial stems from the perceived advantage of using a material similar to the mineral phase in bone and teeth for replacing these materials. Because of this similarity,better tissue bonding is expected. Additional perceived advantages of bioactive ceramics include low thermal and electrical conductivity, elastic properties similar to those of bone, control of in vivo degradation rates through control of material properties, and the possibility of the ceramic functioning as a barrier to metallic corrosion products when it is coated onto a metal substrate.

However, processing-induced phase transformations provoke a considerable change in the in vitro dissolution behavior, and the different struc-

tures and compositions alter the biological reactions. Given the range of chemical compositions available in bioactive ceramics and the resultant fact that pure HA is rarely used, the broader term calcium-phosphate ceramics (CPC) has been proposed in lieu of the more specific term hydroxyapatite. Each individual CPC is then defined by its own unique set of chemical and physical properties.

Mixtures of HA, tricalcium phosphate, and tetracalcium phosphate may evolve as a result of plasma-spraying deposition processes. Physical properties of importance to the functionality of CPCs include powder particle size, particle shape, pore size, pore shape, pore-size distribution, specific surface area, phases present, crystal structure, crystal size, grain size, density, coating thickness, hardness, and surface roughness.

Results of ex-vivo push-out tests indicate that the ceramidmetal bond fails before the ceramic/ tissue bond and is the weak link in the system. Thus the weak ceramidmetal bond and the integrity of that interface over a lengthy service-life of functional loading is reason for concern.

Surface State and Biocompatibility

Implant materials may corrode or wear, leading to the generation of particulate debris, which may in turn elicit both local and systemic biological responses. Metals are more susceptible to electrochemical degradation than are ceramics. Therefore a fundamental criterion for choosing a metallic implant material is that it elicit a minimal biological response. Titanium-based materials are well tolerated by the body because of their passive oxide layers. The main elemental constituents, as well as the minor alloying constituents, can be tolerated by the body in trace amounts. However, larger amounts of metals usually cannot be tolerated. Therefore minimizing mechanical and chemical breakdown of implant materials is a primary objective.

Titanium and other implant metals are in their passive state under typical physiological conditions, and breakdown of passivity should not occur. Both c.p. Ti and Ti-6A1-4V possess excellent corrosion resistance for a full range of oxide

states and pH levels. It is the extremely coherent oxide layer and the fact that titanium repassivates almost instantaneously through surface controlled-oxidation kinetics that renders titanium so corrosion resistant. The low dissolution rate and near chemical inertness of titanium dissolution products allow bone to thrive and therefore osseointegrate with titanium. However, even in its passive condition, titanium is not inert. Titanium ion release that does occur is a result of the chemical dissolution of titanium oxide.

Analyses of implant surfaces are necessary to ensure a twofold requirement. First, implant materials cannot adversely affect local tissues, organ systems, or organ functions. Second, the in vivo environment cannot degrade the implant and compromise its long-term function. The interface zone between an implant and the surrounding tissue is therefore the most important entity in defining the biological response to the implant and the response of the implant to the body.

The success of any implant depends on its bulk and surface properties, the site of implantation, tissue trauma during surgery, and motion at the implant/tissue interface. Surface analysis in implantology therefore aids in material characterization, determining structural and composition changes occurring during processing, identifying biologically induced surface reactions, and analyzing environmental effects on the interfaces.

The surface of a material is most always different in chemical composition, form, and structure than is the bulk material. These differences arise from the molecular arrangement, surface reactions, and contamination. In this regard, the interface chemistry is primarily determined by the properties of the metal oxide and not as much by the metal itself. Little or no similarity is found between the properties of the metal and the properties of the oxide, but the adsorption and desorption phenomena can still be influenced by the properties of the underlying metal. Therefore characterization of surface composition, binding state, and form and function are all important in

the analysis of implant surfaces and implant/ tissue interfaces.

Metallic oxides dictate the type of cellular and protein binding at the implant surface. Surface oxides are continually altered by the inward diffusion of oxygen, and by hydroxide formation and the outward diffusion of metallic ions. Thus a single oxide stoichiometry does not exist.

Summary Although there is no consensus regarding methods of evaluating dental implants and what criteria are most important, clinical evaluations have generally shown that dental implants are successful in about 75% of the cases, 5 years post-implantation. Despite advances in materials' synthesis and processing, surgical technique, and clinical protocols, clinical failures occur at rates of approximately 2% to 5% per year. Causes of failure and current problems with dental implants include: (1) early loosening, stemming from a lack of initial osseointegration; ( 2 ) late loosening, or loss of osseointegration;

(3) bone resorption; (4) infection; ( 5 ) fracture of the implant or abutment; and (6) delamination of the coating from the bulk implant. The most common failure mechanism is alveolar crest resorption, leading to progressive periodontal lesions, decreased areas of supporting tissues, and ultimately implant loosening. Aseptic failures are most often the cumulative result of more than one of the aforementioned factors.

OTHER APPLICATIONS OF WROUGHT TITANIUM

Based on clinical experience with wrought titanium dental implants, wrought titanium crown and bridge applications have been pursued. With the advent of reproducible, high-tolerance machining and processing techniques, such as spark erosion, laser welding and micromachining, and computer-aided design/computer-aided manufacturing, wrought titanium crowns are now possible. Future applications are likely to include partial-denture work, other precision work,

implant-supported restorations, and orthodontic components.

WROUGHT STAINLESS STEEL

Steel is an iron-carbon alloy. The term stainless steel is applied to alloys of iron and carbon that contain chromium, nickel, manganese, and perhaps other metals to improve properties and give the stainless quality to the steel. These alloys differ in composition from the cobalt-chromium, nickel-chromium, and titanium casting alloys. Usually, stainless steel alloys are not cast, hut instead are used in the wrought form in dentistry, which represents the second way that stainless steel differs from the cast base-metal alloys. As a result, the types of appliances formed from these two materials differ. The most common applications of stainless steel for dental purposes at present are in the preparation of orthodontic appliances and fabrication of endodontic instruments, such as files and reamers. Some specialized applications of stainless steel exist for temporary space maintainers, prefabricated crowns, or other appliances placed in the mouth, and for various clinical and laboratory instruments.

COMPOSITION

Several broad classifications of stainless steel are generally recognized. The various groups are referred to asferritic, martensitic, and austenitic, and they have different compositions, properties, and applications. The ferritic stainless steels are chromium steels employed in the manufacture of instruments or equipment parts in which some degree of tarnish resistance is desirable. A wide range of compositions is available in this group, in which amounts of chromium, the principal element contributing to stainless qualities, may vary from 15% to 25%. Elements such as carbon, silicon, and molybdenum are included but are all held within narrow limits.

The martensitic steels also are primarily chromium steels with a lower chromium content

(about 12% to 18%). These steels can be harclened to some degree by heat treatment, and they have a moderate resistance to tarnish. They are used chiefly in the manufacture of instruments and, to a limited degree, for orthodontic appliances.

The austenitic steels represent the alloys used most extensively for dental appliances. The most common austenitic steel used in dentistry is 18- 8 stainless steel, so named because it contains approximately 18% chromium and 8% nickel. The carbon content is between 0.08% and 0.20%, and titanium, manganese, silicon, molybdenum, niobium, and tantalum are present in minor amounts to give important modifications to the properties. The balance (= 72%) is, of course, iron.

FUNCTION OF ALLOYING ELEMENTS AND CHEMICAL RESISTANCE

The corrosion resistance of stainless steel is attributed largely to the presence of chromium in the alloy, and no other element added to iron is as effective in producing resistance to corrosion. Iron cannot be used without chromium additions because iron oxide (Fe,O,), or rust, is not adherent to the bulk metal. Approximately 11% chromium is needed to produce corrosion resistance in pure iron, and the necessary proportion is increased with the addition of carbon to form steel. Chromium resists corrosion well because of the formation of a strongly adherent coating of oxide on the surface, which prevents further reaction with the metal below the surface. The formation of such an oxide layer is called pas- sivation. The surface coating is not visible, even at high magnification, hut the film adds to the metallic luster of the metal. The degree of passivity is influenced by a number of factors, such as alloy composition, heat treatment, surface condition, stress in the appliance, and the environment in which the appliance is placed. In dental applications, the stainless characteristics of the alloys can therefore be altered or lost by excessive heating during assembly or adapta-

tion; using abrasives or reactive cleaning agents, which can alter the surface conditions of the appliance; and even by poor oral hygiene practices over prolonged periods.

Of the stainless steel alloys in general use, the austenitic type of 18-8 stainless steel shows the greatest resistance to corrosion and tarnish. In these alloys, chromium and nickel form solid solutions with the iron, which gives corrosion protection. The chromium composition in these alloys must be between 13% and 28% for optimal corrosion resistance. If the chromium content is less than 13%, the adherent chromium oxide layer does not form. If there is more than 28% chromium, chromium carbides form at the grain boundaries, embrittling the steel. The amount of carbon must also be tightly controlled. If not, carbon will react with chromium, forming these grain-boundary chromium-carbides, which lead to depletion of grain-boundary chromium and decrease corrosion resistance in a process known as sensitization. Molybdenunl increases the resistance to pitting corrosion.

The elements present in small amounts tend to prevent the fornlation of carbides between the carbon present in the alloy and the iron or chromium and, as a result, often are described as stabilizing elements. Some steels, termed stabi- lized stainless steels, contain titanium, niobium, or tantalum, so the carbides that do form are titanium carbides rather than chromium carbides.

The chemical resistance of stainless steel alloys is improved if the surface is clean, smooth, and polished. Irregularities promote electrochemical action on the surface of the alloy. Soldering operations on stainless steel with gold and silver solder may contribute to a reduction in stainless qualities because of electrogalvanic action between dissimilar metals or because of localized, improper composition of the stainless steel wire.

STRESS-RELIEVING TREATMENTS

The 18-8 alloys are not subject to an increase in properties by heat treatment, but they do respond to strain hardening as a result of cold work during adjustment or adaptation of the alloy to

form the appliance. Heat treatment above 650" C results in recrystallization of the microstructure, compositional changes, and formation of chromium-carbides, three factors that can reduce mechanical properties and corrosion resistance.

Appliances formed from these alloys may, however, be subjected to a stress-relieving operation to remove the effects of cold working during fabrication, increase ductility, or produce some degree of hardening with some alloys. If heat treatment is to be performed, it should be held to temperatures between 400" and 500" C for 5 to 120 seconds, depending on the temperature, type of appliance, and alloy being heated. A time of 1 minute at 450" C would represent an average treatment to be used on an orthodontic appliance. Keep in mind that temperatures above 650" C will soften or anneal the alloy, and the properties cannot be restored by further treatment. The main advantage of a low-temperature heat-treating operation is that it establishes a uniformity of properties throughout the appliance after adaptation and fabrication, which may reduce the tendency toward breakage in service. Factors affecting an alloy's ability to be heattreated and stress-relieved include alloy composition, working history (i.e., fabrication procedure), and the duration, temperature, and atmosphere of the heat treatment.

STAINLESS STEEL ORTHODONTIC WIRES

Manipulation Stainless steel wires are processed in the wrought form, through rolling or drawing. Wrought means that the near net shape is formed from the material in a solid state. Stainless steel wires can be formed easily into appliances with special orthodontic pliers and instruments. Once an appliance is fabricated, it should be heat-treated for 1 minute at 450' C to relieve stresses created during fabrication.

The soldering of stainless steel appliances requires skill, and the use of suitable materials is essential. Borax fluxes are not satisfactory, and the use of fluoride-containing flux is required for a successful solder joint. Gold and silver solders may be employed to form the union. Silver solders are considered easier to use than are most

gold solders, and also provide a stronger solder joint. Silver solders also have the advantage of having slightly lower melting temperatures, 600" to 650" C, than most gold solders, which reduces the hazard of overheating the steel during the soldering operation and thus reducing properties.

The strength of a wire in the vicinity of the soldered joint formed with gold solder is lower than that of one formed with silver solder. Many operators therefore use silver solder with stainless steel or rely entirely on spot welding to assemble the appliance. In one study, which used a standard orthodontic blowtorch as a source of heat with a low-fineness gold solder and a fluoride type of flux, the temperature recorded in the center of a joint was 700" C when the least possible amount of heat was used and 800" C when the maximum amount was applied. The temperature was 650" C at 1 mm from the joint and 635" C at 3 mm, but at 5 mm the temperature in the wire was only 440" C. These observations indicate that higher temperatures are being developed during soldering than were believed to exist, and are no doubt responsible for some of the reduction in strength of stainless steel adjacent to the solder joint.

Cleaning and polishing stainless steel appliances are troublesome operations that become necessary after soldering, heat treating, or periods of service in the mouth. The appliance may be pickled in warmed nitric acid, but a gray satin finish will result that requires buffing or mechanical brushing with a fine abrasive to restore the luster of the original material. An electrolytic polishing bath, also know-nas an anodicpolisher,

is useful for restoring the surface appearance to stainless steel appliances, the same as for the cast cobalt-chromium structures mentioned previously.

Properties ANSI/ADA Specification No. 32 for orthodontic wires not containing precious metals describes Type I (low-resilience) and Type I1 (high-resilience) wires. The requirements for flexure yield strength and number of bend cycles are listed in Table 16-3.

Some properties of stainless steel wires when tension, bending, and torsion are applied are listed in Table 16-4 and compared with the properties of nickel-titanium and beta-titanium wires. Of the three types of wires, the stainless steel wire has the highest values of yield strength, elastic modulus, and spring rate and the lowest springback (yield strength/elastic modulus).

Mechanical properties of stainless steel orthodontic wires are listed in Table 16-5for two sizes of wires in as-received and stress-relieved conditions. The higher values of proportional limit and yield strength of the 0.36-mm diameter wire reflect the increased amount of cold-working used to fabricate this size of wire as compared with that required for 0.56-mm diameter wire. The properties of these wires (except tensile strength) are improved by the stress-relieving heat treatment.

Curves of bending moment versus angular deflection are strongly affected by three factors:

(1) the geometry of the wire; (2) the direction of loading, with respect to the wire orientation; and

(3) the thermal history of the wire. Curves of bending moment versus angular deflection are

I

I

1

Property

Proportional limit,l*MF'a

Yield strength, 0.1% offset,+MPa

Tensile strength, MPa

Hardness (Knoop), kg/mm2

Cold-bending, number of 90-degree bends

shown in Fig. 16-5for stainless steel wires having square and rectangular cross-sections.The direction of bending of the rectangular wire has an important effect on the bending moment because the wire is much stiffer in bending in the larger dimension (upper curve versus the middle curve). The curves for the square wire are nearly the same for bending in the direction of the square or the diagonal (bottom curve). The effect

of heating on the bending moment-angular deflection curves for round stainless steel wires is shown in Fig. 16-6. Heating the wire even for a short time (15 seconds) causes a decrease in the stiffness in bending and in the angle at which permanent deformation starts as compared with that of the as-received wire. Increased temperature also leads to recrystallization and reduced properties.

Fig. 16-5 Bending moment-angular deflection curves for square and rectangular 18-8 stainless steel wires. The dimensions of the wires are in mm.

(From Craig RG, editor: Dental materials: a problem-oriented approach, St Louis, 1978, Mosby)

Soldering and spot welding can cause a deterioration in properties if the wire is overheated or underheated. Cross sections of a spot-welded wire are shown in Fig. 16-7. The low setting of the spot welder (underheating) produced an inadequate joint, whereas the high setting (overheating) caused excessive melting and recrystallization of the wrought structure of the wire.

STAINLESS STEEL ENDODONTIC

INSTRUMENTS

Numerous endodontic instruments are classified for hand use or as motor-driven instruments. The most common instruments are the K type of root canal files and reamers. These are manufactured by machining a stainless steel wire into a pyramidal blank, either square or triangular in cross section, and then twisting the blank to form a

spiral cutting edge. A file with a rhombohedral cross section has also been introduced. Examples of a K file, a K reamer, and a rhombohedral file are shown in Fig. 16-8.

Properties Root canals are rarely straight. Therefore, endodontic files need to be able to follow a curved path. As such, bending and torsional properties of files are important. Similar to orthodontic wires, mechanical properties of endodontic files are dependent upon file geometry, direction of loading and material composition. An example of this is illustrated by the bending moment-angular deflection curves of K and rhombohedral endodontic files (sizes 10 to 35) shown in Fig. 16-9. In bending, the rhombohedral files are less stiff than are the K files, and the increase in stiffness for increasing sizes is smaller and more uniform for the rhombohedral files.

Torsional moment-angular rotation curves of K and rhombohedral endodontic files (sizes 20 and 25) are shown for clockwise and counterclockwise rotation in Fig. 16-10.The rhombo-

Angular deflection (degrees)

Fig. 16-9 Bending moment-angular deflection curves of endodontic files, sizes 10 to 35. Vertical marks on curves, Permanent deformation after 90-degree bending.

(From Dolan DW, Craig RG: J Endodont8:260, 1982.)

hedral files are also less stiff than are the K files in torsion, in both directions of rotation. Both types of files fail at lower angular rotation when tested in the counterclockwise direction because counterclockwise rotation increases the tightness of the twist, resulting in brittle fracture, as shown in Fig. 16-11, A. Clockwise rotation causes untwisting of the instrument, thereby allowing more rotation before failing in a ductile manner (Fig. 16-11,B). Operators are therefore cautioned against twisting a file more than a fourth turn in the counterclockwise direction when a file binds in a canal.

ANSI/ADA Specification No. 28 (IS0 3630- 1) for endodontic files and reamers defines geometry, minimum values of torque and angular deflection, and maximum values of stiffness in

D-25

/ D-20

Angular rotation (degrees)

Fig. 16-10 Torsional moment-angular rotation curves of endodontic files, sizes 20 to 25. Clockwise (CW) and counterclockwise (CCW] rotation.

(From Dolan DW, Craig RG: J Endodont8:260, 1982.)

bending for various sizes of files and reamers. Mechanical tests are made in the clockwise direction. No requirements exist for the counterclockwise direction. Stainless steel instruments are required to be resistant to corrosion when tested according to the specification.

A property not covered in the ANSI/ ADA specification,but of importance clinically,is cutting ability. Measurement of cutting ability requires construction of machines that simulate the cutting motion of the instrument. Dentin substrates have been used but produced variation in the data because of biological differences in the hardness of teeth. To overcome this problem, some investigators use acrylic specimens. Cutting ability of individual files varies considerably but is reproducible for the same file with various numbers of strokes. Wear of a file does not appear to affect its cutting ability. Sterilization by dry heat or salt has no effect on the cutting ability of stainless steel files, but autoclave sterilization causes a reduction. Irrigants such as sodium hypochlorite, hydrogen peroxide, and EDTA-urea cause a reduction in cutting ability, whereas a saline irrigant does not cause a reduc-