Orthodontic Wires

Mechanical Properties

Manufacturing of orthodontic wires

Metallic orthodontic wires are manufactured by a series of proprietary steps, typically involving more than one company. Initially, the wire alloy is cast in the form of an ingot, which must be subjected to successive deformation stages until the cross section becomes sufficiently small for wire drawing. Several deformation stages and intermediate heat treatments are required because considerable work hardening of the alloy occurs during wire manufacturing. Important proprietary details include the rate of drawing, the amount of cross-section reduction per pass, the nature of intermediate heat treatments, the die material and lubricant in contact with the wires, and the ambient atmosphere, which would be important for reactive titanium-containing wire alloys. In general, the casting of the starting ingot and the initial mechanical deformation stages are not performed by the orthodontic materials companies that market the wires.


Fig 21-1 Classification of the alloy systems used for metallic orthodontic wires.

Whereas round orthodontic wires are manufactured by drawing through dies, rectangular and square cross-section wires are fabricated from round wires by a rolling process using a Turk’s head, which contains pairs of rolls. The resulting rectangular or square cross-section wires will necessarily have some degree of rounding at the corners, which varies with the specific wire type and the manufacturer. This edge bevel can be of clinical significance for the actual torque delivered by the archwire-bracket combination. Moreover, the surface roughness of the wire, which has a clinically significant effect on the archwire-bracket sliding friction, varies considerably among the various products and is generally greater for the ß-titanium and nickel-titanium wires.

As a result of the wire-drawing sequence, orthodontic wires have a characteristic wrought microstructure. The original equiaxed grain structure of the starting cast ingot is completely eliminated, and when a polished and etched wire specimen is viewed through the optical microscope or the scanning electron microscope, the wrought grain structure appears as a series of closely spaced lines parallel to the original direction of drawing. It is well-known that this microstructure is essential for an orthodontic wire to maintain the desired temper (springiness) or optimal mechanical properties for clinical use. For example, heat treatment of stainless steel wires at temperatures of 700°C (1,300°F) and higher causes rapid softening and loss of the wrought microstructure as a result of recrystallization. Optimum heat treatments involve only the recovery stage of annealing. In practice, clinicians perform heat treatments using an electric resistance (spot) welding apparatus, and the stainless steel, cobalt-chromium-nickel, and some nickel-titanium wires may be advantageously heat treated.

There are four general wire alloys in significant current use: austenitic stainless steel, cobalt-chromium-nickel, ß-titanium, and nickel-titanium. When selecting a wire based on the desired properties for a given patient, it is important to consider the basic alloy composition and the effect that processes, such as drawing, rolling, and heat treatments that are applied by the manufacturer and the clinician, will have on the specific wire properties and archwire-bracket torque delivery characteristics.

In general, an orthodontist should consider the following aspects in the selection of wires: force delivery, elastic working range, ease of joining individual segments to fabricate more complex appliances, and corrosion resistance and biocompatibility in the oral environment. Cost also represents a significant concern for the clinician, and there is a considerable difference between stainless steel archwires and ß-titanium and nickel-titanium arch-wires. However, these more expensive alloys offer unique properties that should be carefully considered when selecting orthodontic wires.

Bending tests and mechanics principles

To understand the mechanical properties of primary importance when comparing different wire types and sizes, it is first necessary to review the underlying principles and terminology. Typically, the mechanical properties of orthodontic wires are determined from some type of bending test, because this mode of deformation is more representative of clinical conditions than the tension test conventionally used for metals. The cantilever bending test in the original form of American National Standards Institute/American Dental Association (ANSI/ADA) specification1 for orthodontic wires not containing precious metals was strongly criticized. The revised specification2 contains a three-point bending test that better simulates clinical interbracket distances and is more suitable for nickel-titanium wires with very low elastic modulus than the previous test, which used the Olsen stiffness tester.

All bending tests involve the measurement of angular or linear deflection of an archwire segment resulting from a bending moment or an applied force. The bending moment or force and the deflection are represented on the vertical and horizontal axes, respectively, of a graphic plot that is generally a straight line at the force levels for elastic deformation of clinical interest. The exception is when very short test spans are used. Although orthodontists typically activate wires somewhat into the permanent deformation range, the bending properties are based on elastic deformation. Although clinical interest is obviously in the unloading characteristics of activated archwires, investigators have generally determined mechanical properties during the initial loading stage. The elastic loading and unloading plots differ for the nickel-titanium wires, although they do not differ for the other alloy types.

There are several basic mechanical properties of orthodontic wires that are determined from the bending test plot:

1. The elastic force delivery is the slope of the initial straight line and is the amount of force required for unit activation. This property determines the stiffness of the wire and is the inverse of the property of flexibility. For those test plots in which bending moment is portrayed as a function of linear or angular deflection, the elastic force delivery is proportional to the slope of the initial straight line.

2. The proportional limit is the highest point on the bending plot at which the force or bending moment remains proportional to the deflection for a wire alloy that exhibits linear elasticity (stainless steel, cobalt-chromium-nickel, and ß-titanium wires). Nickel-titanium wires typically exhibit nonlinear elastic behavior.

3. A force or moment at yielding occurs in response to a designated small amount of permanent bending deflection, and is analogous to the yield strength (YS) for the conventional tension test used to evaluate the mechanical properties of metals.

4. A maximum force or moment will be observed during the bending test, but will not have clinical relevance when substantial permanent deformation has taken place, since only elastic deformation is involved for tooth movement.

5. The maximum elastic activation before the onset of permanent deformation is the value of linear or angular deflection corresponding to the maximum elastic force or moment at the proportional limit. This property is known as the elastic range or working range (sometimes simply termed the range) of the wire.

Wire stiffness or elastic force delivery is dependent on two fundamental factors: (1) the composition and structure of the wire alloy, reflecting both the basic metallurgy and the manufacturing sequence, and (2) the wire segment geometry, that is, the cross-section shape and size and the segment length. The basic metallurgy contribution of the wire alloy is given by the modulus of elasticity (E), or the Young modulus, which relates tensile and compressive elastic stress and strain independent of the specimen cross-section area and length. The elastic modulus values determined from bending and tension tests should be the same, provided that the bending deformation is properly analyzed.

The resistance of a cross-section shape to elastic bending is related to the moment of inertia (I). For a round wire of diameter (d), the moment of inertia is given by:


whereas for a rectangular wire of width (w) and thickness (t) in the plane of bending:


The stiffness of an archwire is inversely proportional to the segment length (l), so that if the length is doubled, the wire flexibility or elastic deflection will be doubled for the same applied force or bending moment. Summarizing these contributions, the elastic stiffness or force delivery characteristics are given by the following expressions:



Many studies in the orthodontic literature use the property of flexural rigidity, given by the product of elastic modulus and moment of inertia, to compare the bending deflections for different wire segments having the same length.

There are two additional useful mechanical properties for orthodontic wires, which are obtained by combining the basic mechanical properties previously discussed (Table 21-1):

1. The modulus of resilience is the area under the elastic force-activation plot and represents the total biomechanical energy per unit volume available for tooth movement when the wire is loaded to the maximum elastic stress or bending moment or, equivalently, unloaded from this level. The modulus of resilience for an orthodontic wire is usually written as (YS)2/2E. The yield strength is generally used to represent the onset of permanent deformation because of the difficulty in precisely locating the proportional limit (PL) on a bending test plot. The formal expression in materials science for the modulus of resilience is (PL)2/2E. It follows that the resilience is much more strongly affected by changes in yield strength or proportional limit than elastic modulus, which is important for the heat-treatment response of cobalt-chromium-nickel and stainless steel wires.

2. The springback for an archwire after unloading is expressed as YS/E, which is approximately equal to the maximum elastic strain or working range of the wire. (The formal expression from materials science for springback would be PL/E.) Since the unloading curve from the permanent deformation range for linearly elastic orthodontic wire alloys (ie, other than nickel-titanium wires) is parallel to the elastic loading curve, the value of YS/E represents the approximate amount of elastic strain released by the archwire on unloading.

Numerous articles listed in the recommended readings have discussed the mechanics of bending tests for orthodontic wires.

Table 21-1 Summary of conventional terminology for important mechanical properties of orthodontic wires*

Rate of force delivery (stiffness) Slope of elastic loading curve or unloading curve Image
Moment at yielding or flexural yield strength Bending moment for designated small amount of permanent deformation (vertical axis of graph) YS
Working range Maximum value of purely elastic deformation (horizontal axis)
Modulus of resilience (resilience) Area under elastic loading curve or unloading curve Image
Springback Elastic strain recovered on unloading from permanent deformation range Image

*The expressions are applicable to tension tests and bending tests for wire alloys exhibiting linear elasticity and sufficiently long specimens.3,4 In these cases, the elastic unloading curve of clinical interest is essentially the same as the initial linear plot for elastic loading. For short loading spans and the nonsuper-elastic nickel-titanium alloys, the unloading curves are nonlinear, and it is difficult to define a modulus of elasticity.
For the idealized definitions, the yield strength (YS) should be replaced in the expressions for modulus of resilience (resilience) and springback by the proportional limit (PL). Other symbols have their usual meanings: E = modulus of elasticity, I = moment of inertia (proportional to d4 for round wire and wt3 for rectangular wire), l = segment length, d = diameter, w = width, t = thickness in plane of bending.

Image Orthodontic Wire Alloys

Stainless steel

Stainless steel wires continue to be popular in clinical orthodontics because of their adequate mechanical properties, good corrosion resistance in the oral environment, and low cost. However, these wires have relatively high values of elastic modulus and thus do not provide the light force delivery that is optimum for orthodontic tooth movement. The wires used in orthodontics are generally American Iron and Steel Institute (AISI) types 302 and 304 austenitic stainless steels, with similar nominal compositions, although the use of 17-7 precipitation-hardening stainless steel was explored. Type 302 is composed of 17% to 19% chromium, 8% to 10% nickel, and 0.15% maximum carbon. Type 304 contains 18% to 20% chromium, 8% to 12% nickel, and 0.08% maximum carbon. The balance of the alloy composition (Table 21-2) is essentially iron (approximately 70%). These are the well-known “18-8” stainless steels, so designated because of the percentages of chromium and nickel in the alloys.

Research has shown that the modulus of elasticity in tension for stainless steel orthodontic wires, where values are more reliable than for bending tests, ranges from about 160 to 180 GPa. These values depend on the manufacturer and temper, and are indicative of differences in alloy compositions, wire drawing procedures, and heat-treatment conditions. For many years, it was not appreciated that the elastic modulus for stainless steel orthodontic wires can be significantly decreased below the 190 to 210 GPa range given in standard physical metallurgy textbooks for annealed stainless steel, although this reduction in elastic modulus was well-known more than five decades ago for heavily cold-worked industrial austenitic stainless steel alloys.

X-ray diffraction has shown that austenitic stainless steel archwires do not necessarily have the single-phase austenitic structure in the as-received condition from the manufacturers. The microstructural phases in these stainless steel wires depend on the manufacturer, temper, and cross-section size; the fundamental factors are the AISI type (particularly carbon content) and thermomechanical processing during manufacturing.

The yield strength for the stainless steel archwires shows a much wider variation than the elastic modulus and has been found to range from approximately 1,100 to 1,500 MPa. After heat treatment, the yield strength can increase to about 1,700 MPa for several wire sizes. Heat treatment also causes significant decreases in residual stress and modest increases (~10%) in resilience. The range in values of mechanical properties in tension for as-received stainless steel wires of clinically important sizes is summarized in Table 21-3. Springback (YS/E) was found to range from 0.0060 to 0.0094 for eight different sizes of as-received stainless steel wires and from 0.0065 to 0.0099 after heat treatment.

Table 21-2 General compositions for four major classes of orthodontic wire alloys

Austenitic stainless steel5 17%–20% Cr, 8%–12% Ni, 0.15% C maximum, balance principally Fe (~ 70%)
Cobalt-chromium-nickel6 (Elgiloy) 40% Co, 20% Cr, 15% Ni, 15.8% Fe, 7% Mo, 2% Mn, 0.16% C, 0.04% Be
ß-titanium (TMA)7 77.8% Ti, 11.3% Mo, 6.6% Zr, 4.3% Sn
Nickel-titanium810 (Nitinol) 55% Ni, 45% Ti (may contain small amounts of Cu or other elements)

Table 21-3 Range of mechanical properties in tension of principal clinical importance for four major orthodontic alloys and as-received wires*

Stainless steel (resilient temper) 160–180 1,100–1,500
Cobalt-chromium-nickel (Elgiloy Blue) 160–190 830–1,000
ß-titanium (TMA) 62–69 690–970
Nickel-titanium (Nitinol) 34 210–410

*The data for modulus of elasticity and yield strength are for round wires with diameters from 0.016 to 0.020 inch and for rectangular wires having cross-section dimensions from 0.017 inch × 0.025 inch to 0.019 inch × 0.025 inch.11,12

1 MPa = 145 psi and 1 GPa = 145,000 psi.

The use of heat treatment to eliminate residual stresses that might cause fracture during manipulation of stainless steel appliances can be important under clinical conditions. However, austenitic stainless steel alloys can be rendered susceptible to intergranular corrosion when heated to temperatures between 400°C and 900°C, due to the formation of chromium carbides at the grain boundaries. These precipitates deplete the amount of chromium near the grain boundaries in the bulk stainless steel below that needed for corrosion resistance. Since the stainless steel alloys must be heated within this temperature range for soldering, clinicians are cautioned to minimize the time required for this process.


Cobalt-chromium-nickel orthodontic wires are similar to stainless steel wires in appearance, mechanical properties (see Table 21-3), and joining characteristics, but have a much different composition and considerably greater heat treatment response. Table 21-2 shows that the most commonly used alloy, Elgiloy (Rocky Mountain Orthodontics) has a complex composition of 40% cobalt, 20% chromium, 15% nickel, 15.8% iron, 7% molybdenum, 2% manganese, 0.15% carbon, and 0.04% beryllium, which is similar to that of some base metal casting alloys for removable partial dentures.

The Elgiloy wires are available in four color-coded tempers: soft, ductile, semi-resilient, and resilient. The differences in mechanical properties arise from proprietary variations in the wire manufacturing process. The soft-temper wires (Elgiloy Blue) are popular because they are easily deformed and shaped into appliances, then heat treated to provide substantially increased values of yield strength and resilience. Increases of 20% to 30% in the yield strength of Elgiloy Blue wires after heat treatment have been reported, and similar heat-treatment responses appear to occur for the soft, ductile, and semi-resilient tempers. The large increases in modulus of resilience arise from the dependence on (YS)2 (see Table 21-1). The effect of heat treatment on mechanical properties has been attributed to complex precipitation processes. Springback for the Elgiloy Blue alloy was found to range from 0.0045 to 0.0065 for five different sizes of as-received wires and from 0.0054 to 0.0074 after heat treatment.

The other Elgiloy tempers are less popular than the soft temper because wires made from these tempers have lower formability and are higher in cost than stainless steel. For Elgiloy Blue, the elastic modulus in tension ranges from about 160 to 190 GPa for as-received wires (see Table 21-3), and from about 180 to 210 GPa after heat treatment. The corresponding ranges in yield strength are approximately 830 to 1,000 MPa in the as-received condition, and 1,100 to 1,400 MPa after heat treatment. It is important to emphasize that the elastic force delivery is nearly the same for stainless steel and Elgiloy Blue archwire segments of the same size and length, as indicated in Table 21-3. A common misconception is that the elastic force delivery is much less for Elgiloy Blue wires compared with stainless steel wires because of the “feel” of the former. It is the yield strength and elastic range that are diminished relative to the more resilient stainless steel wires.


A ß-titanium orthodontic alloy, TMA (Ormco/Sybron), was introduced to the orthodontic profession about 25 years ago. The nominal composition of TMA, which is derived from the two major component elements (titanium-molybdenum alloy), is similar to the 77.8% titanium, 11.3% molybdenum, 6.6% zirconium, and 4.3% tin composition of an industrial Beta III titanium alloy (see Table 21-2). The presence of molybdenum (a ß-stabilizing element) causes the elevated-temperature body-centered cubic ß phase of titanium, rather than the lower-temperature hexagonal close-packed α phase, to be metastable at room temperature. Cold work at ambient temperature or heating to slightly elevated temperature can cause partial transformation of metastable ß-titanium alloys to the α phase. The presence of numerous slip systems for dislocation movement in the ß phase results in excellent formability or capability for permanent deformation. Zirconium and tin contribute increased strength and hardness and hinder formation of an embrittling ω phase during wire processing at elevated temperatures. The TMA alloy has somewhat less than half the elastic force delivery (E ranging from about 62 to 69 GPa) of stainless steel wires, with the yield strength ranging from approximately 690 to 970 MPa (see Table 21-3). Springback for four different sizes of as-received TMA wires was found to range from 0.0094 to 0.011.

Another noteworthy characteristic is that the TMA alloy possesses true weldability. (Welded joints that are fabricated from stainless steel and cobalt-chromium-nickel alloys must be built up with the use of solders to maintain adequate strength.) Optimum conditions for welding TMA with commercial apparatus have been published by Nelson and colleagues.13 Heat treatment by the clinician is not recommended for TMA, although this alloy does respond to a precipitation-hardening procedure. Solution heat treatment between approximately 700°C and 730°C, followed by water quenching, then aging at approximately 480°C, results in precipitation of the α-titanium phase and a peak value for the YS/E ratio.

Research by Kusy and colleagues14 has shown that TMA wires have high surface roughness, which leads to high values of archwire-bracket friction. Ion-implanted TMA wires that have substantially reduced archwire-bracket friction are available from Ormco/Sybron.

With the expiration of the original patent for TMA, new ß-titanium orthodontic wires have been introduced, such as Resolve (Dentsply International) and Beta III Titanium (3M Unitek). Transmission electron microscopic observations of microstructural precipitates in the TMA and Resolve ß-titanium wires suggest that different processing procedures (wire drawing parameters and heat-treatment conditions) are used by the two manufacturers. These differences in processing procedures may have practical significance for the clinically important mechanical properties of ß-titanium wire products. The metallurgy of the ß-titanium wires and other recently introduced titanium alloy wires for orthodontics can be complex, and the interested reader should consult the appropriate textbook references listed at the end of the chapter.

The joining of ß-titanium wires has recently been of substantial research interest. Research by lijima and colleagues15 using a new technique of micro–x-ray diffraction, which is capable of analyzing regions much smaller (50 µm in diameter) than conventional x-ray diffraction, has shown that brazed or soldered wires largely retained the original ß-titanium structure and should be acceptable for clinical use. Further evaluation of mechanical properties and corrosion behavior for the ß-titanium joints is needed.

Development of new ß-titanium alloys and other titanium alloys has been an area of considerable activity by manufacturers in recent years, in part because of the biocompatibility of these nickel-free wires. However, the susceptibility of ß-titanium wires to attack in acidulated fluoride solutions used for mouthwashes and the concomitant degradation of wire properties suggest that future research is needed to address clinical concerns about this potential problem. The susceptibility of ß-titanium wires to hydrogen embrittlement with the possibility of delayed fracture in these acidulated solutions also warrants further investigation.


The fourth wire alloy, nickel-titanium, has also remained a strong focus of materials research as well as considerable marketing activity by manufacturers. The name nitinol, which is a generic name for all of these nickel-titanium alloys, is derived from the words nickel and titanium that make up the composition, along with the acronym for Naval Ordnance Laboratory, which is where these alloys were originally developed by Buehler and associates. The use of nickel-titanium wires for orthodontics originated from work by Andreasen and colleagues in the early 1970s.16,17

Nickel-titanium orthodontic alloys are based on the intermetallic compound NiTi, which has weight percentages of 55% nickel and 45% titanium (see Table 21-2). X-ray energy-dispersive spectroscopic analyses of several nickel-titanium wire products suggest that the compositions are slightly titanium-rich (between 50 and 51 atomic percent titanium).

Although the shape-memory effect associated with engineering nickel-titanium alloys was not available in the original Nitinol wire (3M Unitek), there were two features of considerable importance for clinical orthodontics:

1. The very low elastic modulus (E about 34 GPa in tension) for Nitinol corresponds to about one-fifth of the force delivery for stainless steel archwires and half the force delivery for TMA wires having the same cross-section dimensions and length.

2. Because of the extremely wide elastic working range, 12.5-mm segments in the clinically important size ranges retained a permanent set of not more than 5 degrees after 90-degree cantilever bending by the original ANSI/ADA specification test procedure and release of the applied moment.

The yield strength for the original Nitinol orthodontic wires generally ranges from about 210 to 410 MPa. Springback for six different sizes of as-received wires was found to range from 0.0058 to 0.016.

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May 28, 2016 | Posted by in Dental Materials | Comments Off on Orthodontic Wires
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