Cast and Wrought Base-Metal Alloys

1. The alloy’s chemical nature should not produce harmful toxicologic or allergic effects in the patient or the operator.

2. The chemical properties of the appliance should provide resistance to corrosion and physical changes when in the oral fluids.

3. The physical and mechanical properties, such as thermal conductivity, melting temperature, coefficient of thermal expansion, and strength, should all be satisfactory, meeting certain minimum values and being variable for various appliances.

4. The technical expertise needed for fabrication and use should be feasible for the average dentist and skilled technician.

5. The metals, alloys, and companion materials for fabrication should be plentiful, relatively inexpensive, and readily available, even in periods of emergency.

ANSI/ADA SPECIFICATION NO. 14

According to ANSI/ADA Specification No. 14 (ISO 6871), the weight of chromium should be no less than 20%, and the total weight of chromium, cobalt, and nickel should be no less than 85%. Alloys having other compositions may also be accepted by the ADA, provided the alloys comply satisfactorily with requirements on toxicity, hypersensitivity, and corrosion. Elemental composition to the nearest 0.5% must be marked on the package, along with the presence and percentage of hazardous elements and recommendations for processing the materials. The specification also requires minimum values for elongation (1.5%), yield strength (500 MPa), and elastic modulus (170 GPa).

An important feature of this specification is that it made available a standardized method of testing, which has, in turn, made it possible to compare results from one investigation with those of another.

COMPOSITION

MICROSTRUCTURE OF CAST BASE-METAL ALLOYS

The microstructure of any substance is the basic factor that controls properties. In other words, a change in the physical properties of a material is a strong indication that there must have been some alteration in its microstructure. Sometimes this variation in microstructure cannot be distinguished by ordinary means. Neither cobalt-chromium nor nickel-chromium alloys have simple microstructures, and their microstructures change with slight alterations of manipulative conditions.

The microstructure of cobalt-chromium alloys in the cast condition is inhomogeneous, consisting of an austenitic matrix composed of a solid solution of cobalt and chromium in a cored dendritic structure. The dendritic regions are cobalt-rich, whereas the interdendritic regions can be a quaternary mixture consisting of a cobalt-rich γ-phase; a chromium-rich M₂₃C₆ phase, where M is Co, Cr, or Mo; an M₇C₃ carbide phase; and a chromium- and molybdenum-rich σ-phase. Interdendritic casting porosity is also associated with this structure.

Many elements present in cast base-metal alloys, such as chromium, cobalt, and molybdenum, are carbide-forming elements. Depending on the composition of a cast base-metal alloy and its manipulative condition, it may form many types of carbide. Furthermore, the arrangement of these carbides may also vary depending on the manipulative condition.

The microstructure of a commercial cobaltchromium alloy is illustrated in Fig. 16-1. In Fig. 16-1, A. the carbides are continuous along the grain boundaries. Such a structure is obtained when the metal is cast as soon as it is completely melted. In this condition, the cast alloy possesses low elongation values with a good and clean surface. Carbides that are spherical and discontinuous, like islands, are shown in Fig. 16-1, B. Such a structure can be obtained if the alloy is heated about 100° C above its normal melting temperature; this results in a casting with good elongation values but with a very poor surface because of an increased reaction with the investment. The surface is so poor that the casting cannot be used in dentistry.

Dark eutectoid areas, which are lamellar in nature, are shown in Fig. 16-1, C. Such a structure is responsible for very low elongation values but a good and clean casting. From these three examples, it is clear that microstructure can strongly affect physical and mechanical properties. The microstructure of Ni-Cr alloys is strongly dependent on alloy composition. Alloys containing Be form an interdendritic NiBe phase, as shown in Fig 16-1, D. In fact, during normal metallographic procedures involving acid etching of alloy specimens, the NiBe phase is dissolved; what is seen in the figure is the void area left behind. The susceptibility of the NiBe phase to acid attack has been taken advantage of in developing resin-bonded retainers. The retainer may be etched in selected areas, where a composite-like luting agent can then mechanically adhere to the retainer after curing.

Alloys not containing Be have complicated, multiphase microstructures such as that shown in Figure 16-1, E. The precipitates dispersed within the matrix include complex carbides, and, in alloys where niobium (Nb) is present, Mo-Nb-Si compounds. All these precipitates are relatively unaffected by the short heat-treatment cycles that the alloys are subjected to during the ceramic firing procedures.

HEAT TREATMENT OF BASE-METAL ALLOYS

The early base-metal alloys used in removable partial-denture prostheses were primarily cobalt-chromium and were relatively simple. Heat treating these alloys up to 1 hour at 1000° C did not change their mechanical properties appreciably. Base-metal alloys available today for removable partial-denture prostheses, however, are more complex. Presently, complex cobalt-chromium alloys, nickel-chromium, and iron-chromium alloys are used for this purpose.

Studies have shown that many heat treatments of cobalt-based alloys reduce both the yield strength and elongation. If for any reason some soldering or welding must be performed on these partial dentures, the lowest possible temperature should be used with the shortest possible time of heating to the elevated temperature.

PHYSICAL PROPERTIES

MECHANICAL PROPERTIES

Typical mechanical properties of the partial-denture alloys listed in Table 16-1 have been assembled in Table 16-2, together with a representative range of values for Type IV casting gold alloys subjected to a hardening heat treatment.

CORROSION

Recent research on dental casting alloys has been dominated by studies on corrosion and potential biological effects of metal ion release. In general, in vitro corrosion tests have evaluated a number of important variables, including effects of electrolytic media and artificial saliva, alloy composition, alloy microstructure, and surface state of the metal. These variables account for 2 to 4 orders of magnitude variation in the amount of species released. The surface state of the metal is an extremely important factor influencing corrosion, because the surface composition is almost always different from that of the bulk alloy. Another important consideration is corrosion coupled with wear. Up to three times the mass of metal ions, such as Ni and Be, is released during occlusal rubbing in combination with corrosion than during corrosion alone for Ni-Cr alloys. Nickel-chromium alloys with beryllium fail to satisfy the ISO requirements for release of metal ions. No long-term studies have been performed to monitor the impact of the release of such large concentrations of metal ions on the overall health of patients.

CROWN AND BRIDGE CASTING ALLOYS

The nickel-chromium alloys can be divided into those containing or not containing beryllium. Most of the alloys contain 60% to 80% nickel, 10% to 27% chromium, and 2% to 14% molybdenum. As a comparison, cobalt-chromium alloys contain 53% to 67% cobalt, 25% to 32% chromium, and 2% to 6% molybdenum. Those alloys that contain beryllium contain 1.6% to 2.0% of the element. They may also contain small amounts of aluminum, carbon, cobalt, copper, cerium, gallium, iron, manganese, niobium, silicon, tin, titanium, and zirconium. The low atomic weight of about 9 for beryllium, compared with about 59 for nickel and 52 for chromium, results in atomic percentages of beryllium in these alloys of about 11%. OSHA requirements and recent attempts from standards organizations to lower beryllium content below 0.02% may drastically impact the marketability of beryllium-containing alloys.

Properties of these alloys are similar to those reported in Table 16-2 for the cobalt-chromium alloys. Crown and bridge casting alloys exhibit a higher hardness and elastic modulus than do noble alloys, but they require a slightly different approach in casting and soldering to accommodate their higher solidification shrinkage.

Precautions should be taken to avoid exposure to metallic vapor, dust, or grindings containing beryllium and nickel. The safety standard for pure beryllium dust is 2 μg/m³ of air for a time-weighted, 8-hour day. A higher limit of 25 μg/m³ is allowed for a minimum exposure time of less than 30 minutes. Physiological responses may range from contact dermatitis to severe chemical pneumonitis. Therefore efficient local exhaust and filtration systems as well as adequate general ventilation should be used when casting, finishing, and polishing these beryllium-containing alloys.

The presence of nickel is of greater importance because it is a known allergen. The incidence of allergic sensitivity to nickel has been reported to be from 5 to 10 times higher for females than for males, with 5% to 8% of females showing sensitivity. However, no correlation has been found between the presence of intraoral nickel-based restorations and sensitivity. A cobalt-chromium alloy without nickel or other non-nickel-containing alloy should be used on patients with a medical history indicating an allergic response to nickel. The safety standard for pure nickel is 15 μg/m³ of air for a 40-hour week. To minimize exposure of patients to metallic dust containing nickel or beryllium, intraoral finishing should be done with a high-speed evacuation system and preferably in a wet environment.

OTHER APPLICATIONS OF CAST BASE-METAL ALLOYS

Cast cobalt-chromium alloys serve a useful purpose in appliances other than removable partial-denture restorations. In the surgical repair of bone fractures, alloys of this type have been used for bone plates, screws, various fracture appliances, and splints. Metallic obturators and implants for various purposes can be formed from cast base-metal alloys. The use of cobalt-chromium alloys for surgical purposes is well established, and these alloys have numerous oral surgical uses. They can be implanted directly into the bone structure for long periods without harmful reactions. This favorable response of the tissue is probably attributable to the low solubility and electrogalvanic action of the alloy; the metal is inert and produces no inflammatory response. The product known as surgical Vitallium is used extensively for this purpose. The primary metal used in oral implantology today is titanium (see Chapter 22).

COMMERCIALLY PURE TITANIUM

Commercially pure Ti (CP Ti) is available in four grades, which vary according to the oxygen (0.18 to 0.40 wt%) and iron (0.20 to 0.50 wt%) content. These apparently slight concentration differences have a substantial effect on the physical and mechanical properties.

At room temperature, CP Ti has a hexagonal close-packed (HCP) crystal lattice, which is denoted as the alpha (α) phase. On heating, an allotropic phase transformation occurs. At 883° C, a bodycentered cubic (BCC) phase, which is denoted as the beta (β) phase, forms. A component with predominantly β phase is stronger but more brittle than a component with α-phase microstructure. As with other metals, the temperature and time of processing and heat treatment dictate the amount, ratio, and distribution of phases, overall composition and microstructure, and resultant properties. As a result, casting temperature and cooling procedure are critical factors in ensuring a successful casting.

The density of CP Ti (4.5 g/cm³) is about half the value of many of the other base metals. The modulus (100 GPa) is also about half the value of the other base metals. The yield and ultimate strengths vary, respectively, from 170 to 480 MPa and 240 to 550 MPa, depending on the grade of titanium.

TITANIUM ALLOYS: GENERAL

Alloying elements are added to stabilize either the α or the β phase, by changing the β-transformation temperature. For example, in Ti-6Al-4V, aluminum is an α stabilizer, which expands the α-phase field by increasing the (α + β) to β-transformation temperature, whereas vanadium, as well as copper and palladium, are β stabilizers, which expand the β-phase field by decreasing the (α + β) to β-transformation temperature.

In general, alpha-titanium is weldable, but difficult to form or work with at room temperature. Beta-titanium, however, is malleable at room temperature and is thus used in orthodontics. The (α + β) alloys are strong and formable but difficult to weld. Thermal and thermochemical treatments can refine the postcast microstructures and improve properties.

Ti-6AL-4V

At room temperature, Ti-6Al-4V is a two-phase (α + β) alloy. At about 975° C, an allotropic phase transformation takes place, transforming the microstructure to a single-phase BCC β alloy. Thermal treatments dictate the relative amounts of the α and β phases and the phase morphologies and yield a variety of microstructures and a range of mechanical properties. Microstructural variations depend on whether working and heat treatments were performed above or below the β-transition temperature and on the cooling rate.

Following forging at temperatures in the range of 700° to 950° C, thermal treatments below the β-transition temperature (typically performed at approximately 700° C) produce recrystallized microstructures having fine equiaxed α grains (Fig. 16-2). Equiaxed microstructures are characterized by small (3 to 10 μm), rounded grains that have aspect ratios near unity. This class of microstructure is recommended for Ti-6Al-4V surgical implants.

The mechanical properties of (α + β) titanium alloys are dictated by the amount, size, shape, and morphology of the α phase and the density of α/β interfaces. The tensile and fatigue properties of Ti-6Al-4V have been studied extensively. Microstructures with a small (<20 μm) α-grain size, a well-dispersed β phase, and a small α/β interface area, such as in equiaxed microstructures, resist fatigue crack initiation best and have the best high-cycle fatigue strength (approximately 500 to 700 MPa). Lamellar microstructures, which have a greater α/β surface area and more oriented colonies, have lower fatigue strengths (approximately 300 to 500 MPa) than do equiaxed microstructures.

CAST TITANIUM

Based on the attributes, extensive knowledge, and clinical success of wrought titanium implants, interest has developed in cast titanium for dental applications. Although titanium has been cast for more than 50 years, only recently have nearly precision castings been attainable. For aerospace and medical components, hot isostatic pressing and specific finishing techniques are routinely practiced. However, these techniques are beyond the capabilities and affordability of most dental laboratories.

The two most important factors in casting titanium-based materials are its high melting point (≈1700° C for CP Ti) and chemical reactivity. Because of the high melting point, special melting procedures, cooling cycles, mold material, and casting equipment to prevent metal contamination are required. Titanium readily reacts with gaseous elements such as hydrogen, oxygen, and nitrogen, particularly at high temperatures (>600° C). As a result, any manipulation of titanium at elevated temperatures must be performed in a well-controlled vacuum or inert atmosphere. Without such controls, titanium surfaces will be contaminated with α case, an oxygen-enriched and hardened surface layer, which can be as thick as 100 μm. Surface layers of this thickness reduce strength and ductility and promote cracking because of the embrittling effect of the oxygen. The technology required to overcome these factors is what makes casting titanium relatively more expensive.

Because of the high affinity titanium has for hydrogen, oxygen, and nitrogen, standard crucibles and investment materials cannot be used. Investment materials or face coats for the wax patterns must have oxides that are more stable than the very stable titanium oxide, and must also be able to withstand a temperature sufficient to melt titanium. If this is not the case, oxygen is likely to diffuse into the molten metal. Investment materials using a combination of ZrO₂ type face coat that is backed up a phosphate-bonded silica investment or phosphate investment materials involving inert fillers (ZrO₂, Al₂O₃, MgO) achieve this goal.

Because of the low density of titanium, it is difficult to cast in conventional, centrifugal-force casting machines. In the last 10 to 15 years, advanced casting techniques, which combine centrifugal, vacuum, pressure, and gravity casting, new investment materials, and advanced melting techniques (e.g., electric arc melting) have been developed. These advances have improved the feasibility of casting titanium-based materials in the dental laboratory.

Pure titanium and Ti-alloys such as Ti-6Al-4V have been cast into crowns, removable partial-denture frameworks, and complete denture bases. Titanium alloys have a lower melting point than pure titanium. By alloying titanium, the melting temperature can be lowered to the same temperature as that of nickel-chromium and cobalt-chromium alloys. For example, the Ti-Pd and Ti-Cu alloys have melting points of 1350° C. Lower casting temperatures may also reduce the reactivity of titanium with oxygen and other gases. Other titanium alloys, such as Ti-6Al-4V, Ti-15V, Ti-20Cu, Ti-30Pd, Ti-Co, Ti-Cu, and Ti-Cu-Ni, are still in the experimental stages and have not yet been evaluated in any large clinical studies.

Microstructures of cast titanium materials are similar to those described previously, namely coarse lamellar grains, a result of slow cooling through the β to α or β to (α + β) transformation temperature (Fig. 16-3).

The mechanical properties of cast CP Ti are similar to those of Types III and IV gold alloy, whereas cast Ti-6Al-4V and Ti-15V exhibit properties, except for modulus, similar to those of nickel-chromium and cobalt-chromium alloys.

Recently, cast Ti-6Al-4V microstructures have been refined by temporary alloying with hydrogen. The resulting microstructures (Fig. 16-4) can have α-grain sizes less than 1 μm, aspect ratios near unity, and discontinuous grain-boundary α, microstructural attributes that increase tensile and fatigue strength. These changes in microstructural form and structure result in significant increases in yield strength (974 to 1119 MPa), ultimate strength (1025 to 1152 MPa), and fatigue strength (643 to 669 MPa) as compared with respective values for lamellar (902, 994, and 497 MPa) and equiaxed microstructures (914, 1000, and 590 MPa).

Pure titanium has been cast with a pressure-vacuum casting machine. Other manufacturers have developed centrifugal casting machines that use an electric arc to melt the titanium in an argon or helium atmosphere. Melting is performed in a copper crucible, followed by centrifugal casting into a mold that uses investment. Such machines provide a relatively oxygen-free environment and, with the use of a tungsten arc, can reach temperatures of 2000° C. This latter casting regime has been used to cast CP Ti crowns and full-denture bases. Crowns cast in this manner have been evaluated clinically, and results revealed that, although fitness was inferior to that of silver-palladium alloy, it was superior to that of nickel-chromium. Occlusal adjustment was no more difficult than with conventional crowns, and discoloration, occlusal wear, and plaque retention were similar to other metals. The clinical results of full-denture bases cast in this manner have been acceptable.

Observations of randomly chosen cast crowns using old machines and silica-containing investments have revealed gross surface porosities, to a depth of 75 μm, on both the inside and outside of the surfaces. Mechanical polishing is insufficient to remove this porosity. Internal porosities, sometimes measuring up to 30% of the cross-sectional area, are also readily observed. Surfaces of castings can also be contaminated with α case. The cause of the α case is probably poor atmosphere control or contamination from crucible and mold materials. For optimum functionality of the final casting, the surface layer must be removed during finishing. However, even after the α case is removed, internal oxidation can remain and compromise the mechanical properties of the final appliance. Further examination of such castings has also revealed multiple microcracks at the edges of the margins. Some cracks are as long as 100 μm. Cracks of this length are catastrophic to a notch-sensitive material such as titanium.

As outlined, the difficulties with cast titanium for dental purposes include high melting point and high reactivity, low casting efficiency, inadequate expansion of investment, casting porosity, and difficulty in finishing this metal. From a technical standpoint, titanium is difficult to weld, solder, machine, finish, and adjust. Casting titanium requires expensive equipment.

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.

COMPOSITION

Several broad classifications of stainless steel are generally recognized. The various groups are referred to as ferritic, 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 hardened 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 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 (Fe₂O₃), or rust, is not adherent to the bulk metal. About 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 passivation. The surface coating is not visible, even at high magnification, but 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 adaptation; 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. Molybdenum increases the resistance to pitting corrosion.

The elements present in small amounts tend to prevent the formation 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 stabilized 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 heat-treated 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

STAINLESS STEEL ENDODONTIC INSTRUMENTS

NICKEL-TITANIUM ENDODONTIC INSTRUMENTS

The alloys used in nickel-titanium root canal instruments contain about 56% Ni and 44% Ti by weight, which calculate to be 50% of each by atoms. In some instances, less than 2% of cobalt may be substituted for nickel. Of special interest for this application is that these alloys can change their structure from austenitic (body-centered cubic) to martensitic (hexagonal close-packed) as a function of stress during root canal preparation. When the instrument is used in the canal, the rate of stress levels off due to progressive deformation even if strain is increased due to the transformation to martensite. Note the modulus of Ni-Ti austenite is 120 GPa, and that of martensite is 50 GPa. This effect results in what is termed super-elasticity. When the stress decreases, springback occurs without permanent deformation and a return to the austenitic phase.

The super-elasticity of Ni-Ti permits deformations of 8% strain in endodontic files with complete recovery. This value compares to less than 1% with stainless steel instruments. In addition, the Ni-Ti alloys have higher strengths and lower moduli of elasticity than stainless steel, advantages in preparing curved root canals. Also, it has been shown that Ni-Ti and stainless steel endodontic files do not differ with respect to corrosion resistance.

These improved properties of Ni-Ti root canal instruments have permitted them to be effectively used as engine-driven, rotary instruments. In spite of these improved properties, Ni-Ti instruments can fracture. Studies of cyclic fatigue have shown that the radius of curvature of the file was the most important factor in fatigue resistance, with decreasing radius of curvature (increasing diameter) resulting in decreased fracture time. The fracture was always of a ductile type, suggesting that cyclic fatigue was the major cause of failure.

One drawback of Ni-Ti endodontic instruments is that because of their super-elastic quality they must be manufactured by machining rather than by twisting of tapered wire blanks. The instrument design must therefore be ground into the Ni-Ti tapered blanks, thus increasing their cost.

BASE-METAL PREFABRICATED CROWNS

Stainless steel crowns were introduced in 1950 and are recommended for the permanent restoration of primary teeth, particularly in children suffering from rampant caries or in situations where the crowns of the teeth are destroyed. The approximate composition of stainless steel used for crowns and its mechanical properties are listed in Table 16-6. A titanium-stabilized stainless steel may also be used. For purposes of comparison, the properties of a nickel-based alloy used for permanent prefabricated crowns and tin-based and aluminum-based alloys used for temporary prefabricated crowns are also shown. The mechanical properties of stainless steel and nickel-based materials are similar, and the high ductility is important in the clinical adaptation of the crowns. In addition, they have reasonable hardness and strength and, as a result, are classified as permanent restorations. The tin-based and aluminum-based temporary types also have high ductility, but are soft and have lower yield and tensile strengths and thus do not resist clinical wear as do the stainless steel and nickel-based types.

COMPOSITION

A cobalt-chromium-nickel alloy known as Elgiloy is available in wire and band form for various dental appliances. Elgiloy is nominally composed of 40% cobalt, 20% chromium, 15% nickel, 7% molybdenum, 2% manganese, 0.4% beryllium, 0.15% carbon, 15.4% iron, and 0.05% other. Of particular interest is the fact that beryllium serves to decrease the alloy’s melting point, facilitating manufacturing. In composition, this alloy resembles the cast base-metal alloys more than it does stainless steel.

PROCESSING AND MANIPULATION

The orthodontic wire is supplied in various tempers (amounts of cold work): soft, ductile, semispring temper, and spring temper. The wires are typically formed in a ductile condition, allowing them to be easily deformed and shaped into appliances and then heat-treated to maximize strength. The standard heat treatment, similar to the treatment used to relieve stress in a stainless steel wire, is 482° C for 7 minutes. Low-temperature heat treatment causes a phase change and stress relief. Excessive heat treatment can cause embrittlement.

The fabrication and soldering techniques used with Elgiloy are similar to those used with stainless steel wires. Elgiloy wires should be soldered with a silver solder in the presence of a fluoride flux or joined by spot welding.

PROPERTIES

The properties of Elgiloy as received are similar to those of stainless steel wires; however, its properties can be modified slightly by the heat treatment (7 minutes at 482° C) used to stress relieve a stainless steel wire. The mechanical properties of spring-temper Elgiloy wire are proportional limit, 1610 MPa; 0.2% yield strength, 1930 MPa; tensile strength, 2540 MPa; and Vickers hardness number, 700 kg/mm². The number of 90-degree bends until failure for 0.36-mm and 0.46-mm wires is listed in Table 16-7 for as-received and heat-treated conditions. Note that temper, heat treatment, and wire size all affect this property. Bending moment—angular deflection curves for several Elgiloy wires are shown in Fig. 16-12. The stiffness in bending of the wires is similar; however, the angle at which permanent deformation occurs increases from soft-temper to spring-temper types. The permanent set after a 90-degree bend decreases from soft-temper to spring-temper types. Heat treatment for 7 minutes at 482° C causes an increase in the angle at which permanent deformation occurs but decreases the permanent set.

COMPOSITION AND SHAPE-MEMORY EFFECT

The industrial alloy is 55% nickel and 45% titanium and possesses a temperature transition range (TTR). At temperatures below the TTR, the alloy can be deformed plastically. When the alloy is then heated from below to above the TTR, a temperature-induced crystallographic transformation from a martensitic to an austenitic microstructure occurs and the alloy will return to its original shape. Hence, nickel-titanium is called a shape-memory alloy. The orthodontic alloy contains several percent cobalt to lower the TTR. A number of variations of the Ni-Ti alloy have been developed in dentistry. Compositional variations lead to changes in the martensitic and austenitic start and finish temperatures and mechanical properties. Only those wires with austenitic finish temperatures less than 37° C exhibit super-elasticity.

PROPERTIES AND MANIPULATION

Mechanical properties of an orthodontic nickel-titanium alloy are compared with those of stainless steel and a beta-titanium alloy in tension, bending, and torsion in Table 16-4. The nickel-titanium alloy has the lowest elastic modulus and yield strength but the highest springback (maximum elastic deflection). As shown in Figs. 16-13 and 16-14, nickel-titanium has the lowest spring rate but the highest resiliency in bending and torsion of the three alloys used for orthodontic wires. Clinically, the low elastic modulus and high resiliency mean that lower and more constant forces can be applied with activations and an increased working range. The high springback is important if large deflections are needed, such as with poorly aligned teeth. Nitinol wire requires special bending techniques and cannot be bent over a sharp edge or into a complete loop; thus the wire is more suited for use with pretorqued, preangulated brackets. The alloy is brittle and therefore cannot be soldered or welded, so wires must be joined mechanically.

COMPOSITION AND MICROSTRUCTURE

A titanium-molybdenum alloy known as betatitanium was introduced in 1979 as a wrought orthodontic wire. As discussed previously, cp Ti exists in a hexagonal close-packed crystal lattice at temperatures below 883° C and in a body-centered cubic crystal lattice at higher temperatures. These structures are referred to as alpha-titanium and betatitanium, respectively. The beta form of Ti can be stabilized at room temperature by alloying with certain elements. Beta-titanium alloy for dental use has the composition 78% titanium, 11.5% molybdenum, 6% zirconium, and 4.5% tin and is supplied as wrought wire.

MANIPULATION

Beta-titanium wires can be shaped easily, and the wires can be soldered and welded. Joints can be made by electrical resistance welding. Under proper welding conditions, minimum distortion of the cold-worked microstructure occurs.

PROPERTIES

Compared with stainless steel and Elgiloy wires, beta-titanium wire has lower force magnitudes, a lower elastic modulus, higher springback (maximum elastic deflection), a lower yield strength, and good ductility, weldability, and corrosion resistance. The mechanical properties of betatitanium alloy in tension, bending, and torsion are compared with stainless steel and nickel-titanium alloys in Table 16-4 and Figs. 16-13 and 16-14. Beta-titanium alloy has values of yield strength, modulus of elasticity, and springback intermediate to those of stainless steel and Nitinol. Its formability and weldability are advantages over Nitinol, and it has a larger working range than do stainless steel or Elgiloy wires.

Cast Base-Metal Alloys

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Wrought Base-Metal Alloys

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Titanium

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