Properties and Structures of Orthodontic Wire Materials










477
Properties and
Structures of
Orthodontic
Wire Materials
20
477
CHAPTER
“Design is not just what it looks like and feels like. Design is how it works.”
— Steve Jobs
“Science can amuse and fascinate us all, but it is engineering that changes
the world.”
— Isaac Asimov
Men are only so good as their technical developments allow them to be.”
— George Orwell
OVERVIEW
The design of an orthodontic appliance includes consideration of shape and the amount of
wire, but this chapter discusses an additional important design factor: the material used in the
appliance. Clinically, every appliance could be described as a spring or a series of springs. Springs
have three characteristics: stiffness, maximum force, and range. On a stress-strain level, linear
materials have three properties that relate directly to the above three clinical phenomena: mod-
ulus of elasticity (E), yield strength (YS), and the E/YS ratio. Stainless steel and beta-titanium
have stress-strain curves in the elastic range that are linear. Most of the nickel-titanium (Ni-TI)
wires are not linear and have a more complicated relationship between their clinical properties
and their underlying stress-strain curves. Beta-titanium wires are only about 0.42 the stiffness of
stainless steel. Along with low-stiffness Ni-Ti, variable-modulus orthodontics is possible where
the material is varied rather than the cross section. Superelastic Ni-Ti wires can take advantage
of unique properties, including large elastic deections, relatively constant forces, and thermal
or shape-memory effects. Esthetic wires are typically coated metal wires. However, a newer gen-
eration of esthetic wires is being developed that are clear. Materials used are ber-reinforced
composites and self-reinforced polymers.
A. Jon Goldberg
Charles J. Burstone

20
Properties and Structures of Orthodontic Wire Materials
478
The design of orthodontic appliances continues to
grow in sophistication to achieve more effective
and predictable control of force systems. Force sys-
tems result from a combination of the wire and
bracket design and the wire’s material properties.
The properties of a material are determined by its
composition, atomic structure, and the microstruc-
tural mechanisms responsible for deformation. This
chapter describes the properties of the metal alloys
widely used in orthodontics as well as ber compos-
ites and the future potential for orthodontic wires
based on polymers.
The three primary mechanical characteristics of an
orthodontic wire are range, stiffness, and maximum
force (or moment). Range is the maximum distance
over which the wire can be deected and recover to
apply force. This characteristic is most important in
the early stages of treatment, where larger range
allows engagement of teeth with greater misalign-
ment. The forces generated by a wire are propor-
tional to its stiffness and the amount of deection.
For the early to middle stages of treatment, low
stiffness is desirable because it imparts biological-
ly favorable low and continuous forces. Low stiff-
ness also allows use of wire cross sections that can
ll and fully engage the bracket, allowing for early
three-dimensional control.
During initial deection, the following linear re-
lationship exists among the primary wire properties.
While not exact for all clinical situations, this rela-
tionship is important and useful.
Range Maximum force (or moment)
(maximum deection)
=
Stiffness
In addition to range, stiffness, and maximum
force, there are other important features of a wire.
Formability is important not only for the clinician but
also for the manufacturer to form arch shapes and
customized appliances. Of course, the wires must be
biocompatible and stable in the oral environment.
Finally, there is continued interest in orthodontic
wires that are esthetic. This last feature is the prima-
ry motivation for the introduction of the composite
and polymer-based archwires described at the end
of the chapter.
Mechanical Behavior and
Relationships
Force-deection curves
The mechanical characteristics of an orthodontic
wire are most conveniently represented with one of
two related graphical forms: the force-deection (F/∆)
curve or the stress-strain curve. A force-deection
curve is useful clinically because it illustrates the
force generated with deection of a given wire. F/∆
curves are not meant to be clinical simulations, but
the graph is applicable to various clinical loading
conditions. The curves are typically measured in a
laboratory with simplied conditions, such as load-
ing in the center of a wire supported at its ends
(three-point bending) or loading one end of a wire
that is gripped at the opposite end (free-end canti-
lever). To facilitate comparisons between laborato-
ries, standard methods for testing orthodontic
wires have been developed, such as American Na-
tional Standards Institute/American Dental Associa-
tion (ANSI/ADA) Specication No. 32 and Interna-
tional Standards Organization (ISO) Specication
No. 15841. In addition to the loading conditions, an
F/∆ curve is dependent on the span length and
cross-sectional dimensions of the wire test sample.
For the same wire geometry, the curves are very help-
ful in comparing different orthodontic materials.
A typical F/∆ curve for stainless steel, cobalt-
chromium (Co-Cr; Elgiloy), or beta-titanium (β-Ti) is
shown in Fig 20-1. For all materials, the initial force
imparted by a wire is proportional to the initial
deection, meaning the beginning of the graph is
linear, as indicated by segment O-YP. Materials or
appliances that exhibit force proportional to deec-
tion are said to follow Hooke’s law. The slope of the
linear region (F/∆) is the measure of a wire’s stiffness.
The steeper the slope, the greater the stiffness. For
clinical situations where high stiffness is desirable,
such as when stabilizing the teeth into a segment or
during the later stages of treatment, a steeper slope
is preferred. For initial alignment and leveling stag-
es of treatment, atter slopes that produce lower,
more continuous forces for each unit of deection
are desirable. Any deections in the linear region
are elastic. Accordingly, during unloading due to
tooth movement, the wire will follow the solid YP-O
segment back to the origin. For stainless steel, Co-Cr,
and β-Ti, this is the working region of the wire. The
area under the loading curve is the mechanical en-
ergy applied to the wire, and the area under the un-

479
Mechanical Behavior and Relationships
loading curve is the energy released by the activated
wire. In the elastic range, stored energy is released
without any loss.
As a wire is deected further, it eventually reaches
the end of the linear region (point YP in Fig 20-1).
For stainless steel, Co-Cr, and β-Ti, this is associated
with initiation of the movement of microstructural
features called dislocations (see section later in the
chapter titled “Dislocation-dependent alloys”). The
movement of dislocations is not reversible, so any
deection due to this mechanism is not recoverable;
this is a region of plastic deformation as indicated
in Fig 20-1. Unloading in the region of plastic defor-
mation follows a line parallel to O-YP but intersects
at a distance from the origin, which quanties the
extent of permanent deformation.
The point YP is important because it corresponds
to the amount of force—point a in Fig 20-1—that
the wire can sustain before any permanent defor-
mation occurs. Additionally, the corresponding
point on the deection axis (point b) is a measure
of the amount of elastic deection the wire can sus-
tain, or its range, the distance from O to b.
Referring again to Fig 20-1, as deection increases
beyond point YP, the force continues to increase to
the maximum that the wire can support. This point is
often labeled as the maximum or ultimate force (or
moment). Continual deection beyond the ultimate
force will eventually result in fracture at point x. The
amount of plastic deection that a wire can sustain
(segment b-d) is a measure of the wire’s formabili-
ty. Stainless steel, β-Ti, and Co-Cr (in the appropriate
condition) have large regions of plastic deformation
in their F/∆ curves and therefore are very formable
clinically and can be easily shaped into various con-
gurations.
Stress-strain curves
While F/∆ curves are clinically useful for compar-
ing wires, the graphs are dependent on wire cross
section and length. Normalizing the values facil-
itates analysis of the inherent material properties,
which is useful for both clinical and engineering
evaluations. Each specic appliance conguration
will have a unique F/∆ ratio even if the wire alloy
in different congurations is the same; normalizing
gives universal wire material properties indepen-
dent of design. Dividing force by the cross section
of the wire and dividing deection or deformation
by the original length results in a stress-strain curve.
A typical stress-strain curve is shown in Fig 20-2; its
shape is similar to the F/∆ curve. Stress-strain curves
are usually measured with tensile loading. Stress, of-
ten represented by σ, is a measure of force/area and
can be expressed in MPa (megapascals), psi (pounds/
Fig 20-1 A typical F/∆ curve for stainless steel, Co-Cr, or β-Ti wire.
The stiffness of a wire is indicated by the slope of the linear region,
F/∆. The maximum force in the elastic region is represented by
point YP. Unloading in this region follows the solid blue line back
to the origin, O, with no permanent deformation. Wires loaded
into the region of plastic deformation will follow the dashed blue
line during unloading and sustain permanent deformation. Failure
occurs at point x. The distance from O to b represents the range or
maximum elastic deection of a wire. The distance from b to d is a
measure of the plastic deformation that the wire can sustain, or its
ductility or formability.
Fig 20-2 A typical stress-strain curve for stainless steel, Co-Cr, or
β-Ti wire. These curves represent the inherent material properties
of an alloy. The slope of the linear region is a measure of the mod-
ulus of elasticity (E), the material’s stiffness. Point YP indicates the
initiation of yielding, or the region where plastic deformation be-
gins to occur. It may be measured by the proportional limit, elastic
limit, or offset yield strength. Failure occurs at point x. The distance
from b to d is a measure of the maximum plastic deformation, or
ductility or formability.
Force, F (g)
Elastic
deformation
Plastic
deformation
Maximum
YP
F/∆
Stiffness
Deection, ∆ (mm)
a
O
b
d
O
b
d
Strain, ε (mm/mm)
Stress, σ
(MPa)
x x
Elastic
deformation
Plastic
deformation
Ultimate strength
Proportional limit
YP
E
0.2%
Elastic limit
Offset yield
strength

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477Properties and Structures of Orthodontic Wire Materials 20477CHAPTER“Design is not just what it looks like and feels like. Design is how it works.” — Steve Jobs“Science can amuse and fascinate us all, but it is engineering that changes the world.” — Isaac Asimov“Men are only so good as their technical developments allow them to be.” — George OrwellOVERVIEWThe design of an orthodontic appliance includes consideration of shape and the amount of wire, but this chapter discusses an additional important design factor: the material used in the appliance. Clinically, every appliance could be described as a spring or a series of springs. Springs have three characteristics: stiffness, maximum force, and range. On a stress-strain level, linear materials have three properties that relate directly to the above three clinical phenomena: mod-ulus of elasticity (E), yield strength (YS), and the E/YS ratio. Stainless steel and beta-titanium have stress-strain curves in the elastic range that are linear. Most of the nickel-titanium (Ni-TI) wires are not linear and have a more complicated relationship between their clinical properties and their underlying stress-strain curves. Beta-titanium wires are only about 0.42 the stiffness of stainless steel. Along with low-stiffness Ni-Ti, variable-modulus orthodontics is possible where the material is varied rather than the cross section. Superelastic Ni-Ti wires can take advantage of unique properties, including large elastic deections, relatively constant forces, and thermal or shape-memory effects. Esthetic wires are typically coated metal wires. However, a newer gen-eration of esthetic wires is being developed that are clear. Materials used are ber-reinforced composites and self-reinforced polymers.A. Jon GoldbergCharles J. Burstone 20Properties and Structures of Orthodontic Wire Materials478The design of orthodontic appliances continues to grow in sophistication to achieve more effective and predictable control of force systems. Force sys-tems result from a combination of the wire and bracket design and the wire’s material properties. The properties of a material are determined by its composition, atomic structure, and the microstruc-tural mechanisms responsible for deformation. This chapter describes the properties of the metal alloys widely used in orthodontics as well as ber compos-ites and the future potential for orthodontic wires based on polymers.The three primary mechanical characteristics of an orthodontic wire are range, stiffness, and maximum force (or moment). Range is the maximum distance over which the wire can be deected and recover to apply force. This characteristic is most important in the early stages of treatment, where larger range allows engagement of teeth with greater misalign-ment. The forces generated by a wire are propor-tional to its stiffness and the amount of deection. For the early to middle stages of treatment, low stiffness is desirable because it imparts biological-ly favorable low and continuous forces. Low stiff-ness also allows use of wire cross sections that can ll and fully engage the bracket, allowing for early three-dimensional control. During initial deection, the following linear re-lationship exists among the primary wire properties. While not exact for all clinical situations, this rela-tionship is important and useful. Range Maximum force (or moment) (maximum deection) = StiffnessIn addition to range, stiffness, and maximum force, there are other important features of a wire. Formability is important not only for the clinician but also for the manufacturer to form arch shapes and customized appliances. Of course, the wires must be biocompatible and stable in the oral environment. Finally, there is continued interest in orthodontic wires that are esthetic. This last feature is the prima-ry motivation for the introduction of the composite and polymer-based archwires described at the end of the chapter.Mechanical Behavior and RelationshipsForce-deection curvesThe mechanical characteristics of an orthodontic wire are most conveniently represented with one of two related graphical forms: the force-deection (F/∆) curve or the stress-strain curve. A force-deection curve is useful clinically because it illustrates the force generated with deection of a given wire. F/∆ curves are not meant to be clinical simulations, but the graph is applicable to various clinical loading conditions. The curves are typically measured in a laboratory with simplied conditions, such as load-ing in the center of a wire supported at its ends (three-point bending) or loading one end of a wire that is gripped at the opposite end (free-end canti-lever). To facilitate comparisons between laborato-ries, standard methods for testing orthodontic wires have been developed, such as American Na-tional Standards Institute/American Dental Associa-tion (ANSI/ADA) Specication No. 32 and Interna-tional Standards Organization (ISO) Specication No. 15841. In addition to the loading conditions, an F/∆ curve is dependent on the span length and cross-sectional dimensions of the wire test sample. For the same wire geometry, the curves are very help-ful in comparing different orthodontic materials.A typical F/∆ curve for stainless steel, cobalt- chromium (Co-Cr; Elgiloy), or beta-titanium (β-Ti) is shown in Fig 20-1. For all materials, the initial force imparted by a wire is proportional to the initial deection, meaning the beginning of the graph is linear, as indicated by segment O-YP. Materials or appliances that exhibit force proportional to deec-tion are said to follow Hooke’s law. The slope of the linear region (F/∆) is the measure of a wire’s stiffness. The steeper the slope, the greater the stiffness. For clinical situations where high stiffness is desirable, such as when stabilizing the teeth into a segment or during the later stages of treatment, a steeper slope is preferred. For initial alignment and leveling stag-es of treatment, atter slopes that produce lower, more continuous forces for each unit of deection are desirable. Any deections in the linear region are elastic. Accordingly, during unloading due to tooth movement, the wire will follow the solid YP-O segment back to the origin. For stainless steel, Co-Cr, and β-Ti, this is the working region of the wire. The area under the loading curve is the mechanical en-ergy applied to the wire, and the area under the un- 479Mechanical Behavior and Relationshipsloading curve is the energy released by the activated wire. In the elastic range, stored energy is released without any loss.As a wire is deected further, it eventually reaches the end of the linear region (point YP in Fig 20-1). For stainless steel, Co-Cr, and β-Ti, this is associated with initiation of the movement of microstructural features called dislocations (see section later in the chapter titled “Dislocation-dependent alloys”). The movement of dislocations is not reversible, so any deection due to this mechanism is not recoverable; this is a region of plastic deformation as indicated in Fig 20-1. Unloading in the region of plastic defor-mation follows a line parallel to O-YP but intersects at a distance from the origin, which quanties the extent of permanent deformation. The point YP is important because it corresponds to the amount of force—point a in Fig 20-1—that the wire can sustain before any permanent defor-mation occurs. Additionally, the corresponding point on the deection axis (point b) is a measure of the amount of elastic deection the wire can sus-tain, or its range, the distance from O to b. Referring again to Fig 20-1, as deection increases beyond point YP, the force continues to increase to the maximum that the wire can support. This point is often labeled as the maximum or ultimate force (or moment). Continual deection beyond the ultimate force will eventually result in fracture at point x. The amount of plastic deection that a wire can sustain (segment b-d) is a measure of the wire’s formabili-ty. Stainless steel, β-Ti, and Co-Cr (in the appropriate condition) have large regions of plastic deformation in their F/∆ curves and therefore are very formable clinically and can be easily shaped into various con-gurations. Stress-strain curvesWhile F/∆ curves are clinically useful for compar-ing wires, the graphs are dependent on wire cross section and length. Normalizing the values facil-itates analysis of the inherent material properties, which is useful for both clinical and engineering evaluations. Each specic appliance conguration will have a unique F/∆ ratio even if the wire alloy in different congurations is the same; normalizing gives universal wire material properties indepen-dent of design. Dividing force by the cross section of the wire and dividing deection or deformation by the original length results in a stress-strain curve. A typical stress-strain curve is shown in Fig 20-2; its shape is similar to the F/∆ curve. Stress-strain curves are usually measured with tensile loading. Stress, of-ten represented by σ, is a measure of force/area and can be expressed in MPa (megapascals), psi (pounds/Fig 20-1 A typical F/∆ curve for stainless steel, Co-Cr, or β-Ti wire. The stiffness of a wire is indicated by the slope of the linear region, F/∆. The maximum force in the elastic region is represented by point YP. Unloading in this region follows the solid blue line back to the origin, O, with no permanent deformation. Wires loaded into the region of plastic deformation will follow the dashed blue line during unloading and sustain permanent deformation. Failure occurs at point x. The distance from O to b represents the range or maximum elastic deection of a wire. The distance from b to d is a measure of the plastic deformation that the wire can sustain, or its ductility or formability.Fig 20-2 A typical stress-strain curve for stainless steel, Co-Cr, or β-Ti wire. These curves represent the inherent material properties of an alloy. The slope of the linear region is a measure of the mod-ulus of elasticity (E), the material’s stiffness. Point YP indicates the initiation of yielding, or the region where plastic deformation be-gins to occur. It may be measured by the proportional limit, elastic limit, or offset yield strength. Failure occurs at point x. The distance from b to d is a measure of the maximum plastic deformation, or ductility or formability.Force, F (g)ElasticdeformationPlasticdeformationMaximumYPF/∆StiffnessDeection, ∆ (mm)aObdObdStrain, ε (mm/mm)Stress, σ(MPa)x xElasticdeformationPlasticdeformationUltimate strengthProportional limitYPE0.2%Elastic limitOffset yield strength 20Properties and Structures of Orthodontic Wire Materials480square inch), or other comparable units. Strain (ε) is a measure of the deformation per original size of the sample and is expressed as mm/mm, inch/inch, or dimensionless units. Because of its importance, the slope of the initial linear region of a stress-strain curve has a specic name: the modulus of elasticity or the Young modulus, which is abbreviated as E. In a stress-strain curve, the point at which the wire changes from elastic to plastic deformation is called the yield point or yield stress and is designated YP or YS, respectively. There are different denitions of where on the curve the transition from linear to nonlinear behavior occurs. The point at which the line changes from linear to nonlinear is the propor-tional limit. The precise point at which the materi-al transitions from elastic to plastic behavior is the elastic limit. Because it is difcult to experimentally detect these values precisely, a common method is to construct a line parallel to the linear region but offset by a predetermined amount of typically 0.2%. The intersection of this line with the stress-strain curve is the offset yield strength. In the stress-strain curve shown in Fig 20-2, the distance from O to b is equal to the ratio of YS/E. This is a useful relation-ship because it explains not only why a lower mod-ulus imparts lower forces for a given deection but also why for a given YS, a lower E increases working range (O-b).The area under the linear region of the graph, or the area enclosed by points O, YP, and b, is a mea-sure of the resilience or elastic energy stored in a wire when deected to the YP. It is sometimes used to indicate the elasticity of a wire, although the ratio YS/E is a more clinically relevant measure of the working range. The total area under the stress-strain curve is a measure of a material’s toughness. Extremely high-strength wires will have a high YP but limited plastic deformation (segment b-d), re-sulting in a lower toughness, which is the quantita-tive measure of brittleness.The F/∆ and stress-strain curves, and all of the de-rived properties, are based on test methods where the sample wire is loaded continuously until failure. However, in clinical use, appliances are exposed to multiple cycles of force or stress below the YP. This loading condition can result in a cumulative effect that causes failure, even if the maximum load is nev-er exceeded. This is known as fatigue.The modulus of elasticity is an inherent physical property of the material that is not changed by physical stimulus such as bending, torsion, or even heat treatment. Annealing with high temperature will reduce the YP but not the modulus.Relationships between material properties and exure behaviorThe similarity between stress-strain curves and F/∆ curves can be quantied with very important and useful relationships between the engineering mate-rial properties and the clinical exure behavior of orthodontic wires. While the following formulas are strictly correct only with small deections and with-in the initial linear regions of the curves, they can estimate beyond these conditions and illustrate the dependence of orthodontic force systems on mate-rial properties and wire geometry. The stiffness of a wire in exure, or its F/∆ rate in a F/∆ curve, is related to the modulus of elasticity (E) of the material and the wire geometry by the following relationship:F/∆ = EIKL3where I is the moment of inertia and is related to the cross-sectional size, shape, and direction of bending; K is a geometric factor related to the conguration and loading conditions; and L is the wire length. The maximum bending moment (Mmax) that a wire can generate is calculated by the following: Mmax = YS C/Iwhere YS is the yield strength, C is the cross-section radius, and I is the moment of inertia. Because I var-ies exponentially with cross-sectional dimensions and therefore has a signicant effect on appliance stiffness, it is convenient to dene a cross-section stiffness number. These values provide a convenient comparison of wires or appliances of different wire diameters, relative to a clinically useful baseline. Cross-section stiffness numbers and the related ma-terial stiffness numbers are described in chapter 21, where useful tables of these values are also provided.Crystal Structure and Phase TransitionsAll metal alloys used in orthodontics are crystalline, consisting of very specic arrangements of atoms. An example of one atomic pattern that is possible in Ni-Ti alloys is shown in Fig 20-3a. This pattern is referred to as body-centered cubic (BCC). It is a square arrangement with an atom in each corner 481Composition and Properties of Orthodontic Alloysforming a cubic shape and an atom in the center of the cube. Another arrangement known as end- centered monoclinic is shown in Fig 20-3b. Schematic diagrams showing the patterns of crystals are called unit cells. Other atomic arrangements are possible and involve changing the lengths or relative angles between the axes of the unit cell or different po-sitioning of interior atoms, such as on the base or the faces of the unit cell. In nature, there are only 14 possible arrangements, referred to as the Bravais lattices. All crystalline materials adopt one of these specic lattices, but many materials, such as memory alloys, can transition between different lattices with temperature, processing conditions in the manufac-ture of the wire, or stress during clinical application. The different lattice arrangements of an alloy can also be referred to as phases. The intrinsic mechani-cal properties of an orthodontic wire are dependent on both its composition and the phases that are present. Some alloy phases are particularly import-ant and are assigned names. For example, the phase that is stable at higher temperatures is sometimes referred to as austenite or the austenitic phase. Most phase changes require diffusion of the atoms over distances of many unit cells. However, there are phase transitions that result from stress or tempera-ture changes that cause shifts of only one atomic distance to adjacent positions in the lattice. These are often referred to as martensitic or the marten-site phase. Importantly for orthodontics, because the atoms in the martensite phase are so close to their position in the previous phase, the atoms can shift back and thereby appear to remember their original position. Contemporary orthodontic alloys make use of different types of phase transitions.Composition and Properties of Orthodontic AlloysThe four broad categories of metal alloys used in orthodontics include stainless steel, Co-Cr, β-Ti, and Ni-Ti. In addition to composition and phase or lattice-pattern arrangements, it is useful to discuss the alloys according to the atomic mechanisms re-sponsible for their unique mechanical characteristics. The properties of the rst three categories of alloys are dependent on microstructural features called dislocations. Most of the Ni-Ti alloys derive their me-chanical characteristics from phase transitions. The compositions and mechanical characteristics of the various alloys are described below, followed by an ex-planation of the dislocation-dependent and phase transition–dependent microstructural mechanisms.Dislocation-dependent alloysStainless steelIn the mid-1900s, stainless steel replaced gold-based alloys as the most widely used orthodontic wire al-loy. Today, it remains one of the standard alloys for clinical use because of its overall mechanical proper-ties, low coefcient of friction, and low cost.Stainless steel is iron alloyed with chromium, nick-el, and less than 1% carbon. As with all orthodon-tic alloys, the properties are dependent on both composition and atomic structure, the latter being determined by the processing to form the wire. Or-thodontic stainless steels are most commonly AISI Fig 20-3 Examples of two atomic lattice arrangements that can be adopted by metals: (a) body-centered cubic (BCC) arrangement; (b) end-centered monoclinic arrangement.a b 20Properties and Structures of Orthodontic Wire Materials482304 series alloys containing 18% chromium and 8% nickel and are sometimes referred to as 18-8 stain-less. The chromium imparts corrosion resistance and along with the nickel stabilizes the austenitic atomic lattice structure. The high mechanical properties of stainless steel are primarily due to the cold working that occurs during drawing of the wire to clinically relevant cross-sectional dimensions. High amounts of cold working produce wires with very high strength and high elastic springback, but formabili-ty may be limited. Therefore, most commercial wires have a history of only moderate work hardening to minimize brittleness. Figure 20-2 is representative of the shape of the stress-strain curve for stainless steel. The stiffness or modulus (E) of stainless steel is approximately 180 GPa, the highest of all the orthodontic alloys and comparable to that of Co-Cr. Depending on the amount of cold reduction during wire drawing, the YS can vary from 1,200 to 1,930 MPa, with a corresponding decrease in formability. The mechanical property values are summarized in Table 20-1.Cobalt-chromiumCo-Cr orthodontic alloys have a long history of use in orthodontic therapy. Their unique feature is the ability to be easily formed into a desired shape while in a softened condition, then strengthened with a short, in-ofce heat treatment to develop properties more effective for force application. The alloy, originally developed for watch springs, was in-troduced as Elgiloy, but other brands are now avail-able, such as Colboloy (G&H Orthodontics).The wires are available in several starting, soft-ened conditions that can be readily formed. After forming to the desired shape, the wires are strength-ened with heat treatments lasting typically 5 to 10 minutes at 480oC (896oF). The heat treatments in-crease YS and springback but decrease formability, with the extent of the effects dependent on the starting condition of the wire. Heat treating does not alter the modulus of elasticity, so the forces for any given activation are not changed. Representa-tive values of the mechanical properties are shown in Table 20-1. After heat treatment, the exure properties of Co-Cr wires are comparable to those of stainless steel. The ability to have both good form-ability and high strength is useful, but Co-Cr wires are less popular now because after heat treatment their properties are comparable to those of stainless steel, preformed arches are readily available, and heat treatment involves an added step.Beta-titaniumTitanium has been an important structural metal in various industries for over 50 years because of its high ratio of strength to weight, corrosion resis-tance, and biocompatibility. The primary industrial titanium is alloyed with 6% aluminum and 4% va-nadium. In dentistry, the success of implants is due to the ability of bone to grow against and maintain intimate contact with commercially pure (99%) ti-tanium. The structural and implant applications use titanium that is partially or completely in its hex-agonal close-packed or alpha lattice arrangement. However, as an orthodontic archwire, the YS and elastic range of alpha titanium provides little ben-et over stainless steel. However, with the addition of molybdenum, titanium acquires the BCC lattice or β-Ti structure. The modulus of β-Ti is about 35% less than the alpha form of titanium and about 42% of that of stainless steel and Co-Cr orthodontic alloys. Accordingly, for the same deection, the forces de-livered by a β-Ti wire are 42% of those of stainless Table 20-1 Mechanical properties of orthodontic wire materialsEngineering term Modulus, E (GPa) Yield strength, YS (MPa)YS/E (×10–3)Clinical term Stiffness Strength Range FormabilityStainless steel 159–200 1,200–1,930* 8.69 Low–high*Co-Cr 150–211 1,400 7.78 High†β-Ti68–72 960–1,170 15.4 HighSuperelastic Ni-Ti See F/∆ curve‡450–600 Very high§Not formableMartensitic Ni-Ti 33||1,655 50.2 Not formable*Strength and formability depend on the amount of cold working during wire drawing. More cold working increases strength but decreases formability.†Formability of Co-Cr alloys is high in the softened, non–heat-treated condition.‡Modulus varies with activation (see Fig 20-6).§Range is very high and varies with activation.|| Modulus of the initial linear region of the stress-strain curve. 483Composition and Properties of Orthodontic Alloyssteel or Co-Cr wires of the same dimension. With proper cold working to increase strength, the elas-tic range of β-Ti becomes about 50% greater than that of stainless steel yet retains comparable form-ability. Overall, the balance of a low modulus, high strength, and good formability makes β-Ti a useful orthodontic archwire alloy. β-Ti was originally intro-duced as titanium-molybdenum alloy (TMA), but other brands are now also available (Beta III, Unitek; BT3, G&H Orthodontics). The mechanical properties of β-Ti are summarized in Table 20-1.Martensitic and austenitic nickel-titaniumNi-Ti orthodontic alloys are important and unique because they deliver exceptionally low forces over very large displacements. There are different types of Ni-Ti alloys used in orthodontic therapy and a few different approaches for categorizing them. Here the Ni-Ti wires are divided into two categories on the basis of the prominent atomic phase during clin-ical use: martensite and austenite. This approach is useful for distinguishing both the mechanical prop-erties and the microstructural mechanisms responsi-ble for their unique performance.The martensitic products were introduced to ortho-dontics in the 1970s and are associated with the trade name Nitinol (Unitek). They are approximately 50% nickel by atomic percentage (55% by weight), with the balance being titanium. Their initial popularity was due to their low stiffness and large elastic range compared with that of stainless steel. A typical F/∆ curve for martensitic Ni-Ti is shown in Fig 20-4, and the mechanical properties are listed in Table 20-1. The austenitic Ni-Ti wires were introduced in the 1980s. These wires produce lower forces and a unique plateau region in their F/∆ curve that can provide almost constant force over large displace-ments, as shown in Fig 20-4. This plateau region is referred to as superelasticity or pseudoelasticity, and the microstructural mechanism responsible for this behavior is described later in the section titled “Superelasticity and phase transitions.”In addition to superelasticity, Ni-Ti wires can pro-duce shape-memory effects (see later section titled “Shape memory”). The earlier Nitinol wires were in the martensitic phase at mouth temperature and had high springback because of work hardening. They became shape-memory alloys at very high temperatures far beyond mouth temperature. The austenitic Ni-Ti wires have largely replaced the mar-tensitic products and are most popular for the ini-tial stages of therapy, when the teeth are at their greatest discrepancy; they take advantage of both superelasticity and shape memory.Phase transition–dependent alloys: Superelastic austenitic Ni-Ti alloysRelative to the martensitic Ni-Ti alloys, the super-elastic austenitic Ni-Ti wires have a slightly different ratio of nickel to titanium or may contain several percent copper, cobalt, or chromium substituting for the nickel. With activation, these alloys trans-form from austenitic to martensitic, and the transi-tion reverses during unloading. This phase change allows these wires to undergo large deformations at almost constant load, indicated by segment a-b in Fig 20-4.While all of the austenitic Ni-Ti wires produce low forces relative to stainless steel, β-Ti, and Co-Cr alloys, there are wide differences among austen-itic Ni-Ti products due to composition, processing, temperature, and particularly clinical loading con-ditions. For example, a complete stress-strain curve (comparable in shape to its F/∆ curve) for austenitic Ni-Ti is shown in Fig 20-5. The initial linear segment of the curve (O-a) has a steeper slope, so the wire is stiffer (higher load-deection rate) with low deec-Fig 20-4 Typical shape of the F/∆ curves for martensitic and aus-tenitic Ni-Ti alloys. Martensitic Ni-Ti does not undergo any phase transitions during activation (solid line) and deactivation (dashed line). Austenitic Ni-Ti does experience a phase transition. Between points a and b, the wire transforms from the austenitic to the mar-tensitic structure, and the transition is reversed during unloading. This mechanical behavior is referred to as superelasticity.LoadDeectionMartensiticAusteniticbaO 20Properties and Structures of Orthodontic Wire Materials484tions. In the plateau region (a-b), there is almost a constant load-deection rate over a large range of deections. At point b, the transition from austen-ite to martensite is complete, but the wire can be loaded further in the all-martensitic state until fail-ure at point c. Unloading follows a parallel path (the dashed line), with the plateau at a lower stress or force. The different unloading path is due to energy loss and is referred to as hysteresis. The amount of applied force, which is the unloading portion of the curve, is also dependent on the extent of activation, as shown in Fig 20-6. A representative range of the mechanical properties for austenitic Ni-Ti products is provided in Table 20-1. While the Ni-Ti wires deliv-er favorable low forces over large displacements, in practical use these wires cannot be formed at room or mouth temperature. Microstructural Mechanisms That Determine Mechanical PropertiesThe previous sections of this chapter have described the composition, atomic lattice structure, and me-chanical properties of stainless steel, Co-Cr, β-Ti, and Ni-Ti orthodontic alloys. In this section, we describe the mechanisms, or the atomic movements during wire activation and forming, that dictate mechani-cal behavior.For all alloys, the initial deection of a wire is due to elastic stretching of the bonds between the at-oms. If the force is removed, the bonds completely recover because the atoms have not been displaced from their lattice positions. This bond-stretching mechanism is responsible for the initial linear region in F/∆ or stress-strain curves. If the force exceeds the bond strength, the atoms are displaced from their equilibrium positions and follow one of two mecha-nisms described below.Dislocation-dependent alloysWhile most of an alloy’s crystal structure is the regu-lar specic spatial arrangement of the atoms, there are also defects in the pattern. One of the most important defects is a partial or incomplete plane of atoms, referred to as a dislocation (Fig 20-7). As stainless steel, Co-Cr, and ß-Ti wires are deect-ed, the stress eventually reaches a level (the yield stress) where the dislocations begin to move with-in the crystalline array of atoms. At this point, the wire does not fracture. The cumulative effect of the movement of millions of dislocations is what allows the wire to plastically deform. The key factor that controls the mechanical properties of these alloys is the mobility of the dislocations. If dislocations can easily move relatively large distances through the crystal structure, the alloy will be formable and have a lower YS. If the movement of dislocations is inhib-ited or blocked, then the wire will be less formable but will have higher strength. Three microstructur-al features can restrict dislocation movement, grain boundaries, second phases, and higher densities of dislocations. Increasing the amounts of these fea-tures will strengthen an alloy but simultaneously Fig 20-5 Complete austenitic Ni-Ti stress-strain (F/∆) curve. Seg-ment O-a is the elastic behavior in the austenite phase. Section a-b is the superelastic or pseudoelastic region where austenite trans-forms to martensite. At point b, the transformation to martensite is complete, and further loading will result in failure at point c. Unloading follows the dashed line and produces low, almost con-stant forces.Fig 20-6 Bending moment versus deection curve for superelastic austenitic Ni-Ti, illustrating that the force applied during unloading is dependent on the amount of activation. (Adapted from Burstone et al1 with permission.)bcaOStrainStressAustenite MartensiteAustenite MartensiteDeection (degrees)Bending moment (gmm) 485Microstructural Mechanisms That Determine Mechanical Propertiesdecrease formability. This explains the effects of commonly employed wire-manufacturing and ma-nipulative procedures, as described below.During the drawing process, reduction of a wire’s diameter down to clinically relevant dimensions in-creases the number of dislocations. The increased dislocation density causes entanglements, inhibit-ing the movement of all dislocations and increasing strength. Accordingly, cold working or work hard-ening is an important factor in developing high strength in orthodontic wires. However, because formability is also dependent on dislocation move-ment, the increased density decreases formability. This explains why the very highest-strength stainless steel wires have limited formability.The presence of a second phase acts like a road-block to the dislocations. The controllable formation of a second phase in the Co-Cr orthodontic alloys, such as Elgiloy, gives this class of wires its unique me-chanical characteristics. The manufacturer supplies wires in a condition that allows signicant formabil-ity. After forming, the clinician heat-treats the wire, generating a second phase that increases strength to levels comparable to those of stainless steel. The heat treatment and precipitation of a second phase does not alter the modulus of elasticity, so the stiff-ness of the wire remains the same.Phase transition–dependent alloysSuperelasticity and phase transitionsAs with other alloys, the initial response of super-elastic austenitic Ni-Ti to deection or load is elas-tic stretching of atomic bonds. When the load is removed, the bonds recover, producing the linear, elastic region O-a in Fig 20-5. If the force is increased beyond point a, instead of dislocation movement the crystal structure undergoes a phase transition.As explained earlier in this chapter, the crystal structures of metals are very specic, and only 14 crystal patterns exist. Many alloys can adopt more than one lattice type, and generally this pattern is determined by the processing conditions during manufacture and the temperature and state of stress during clinical use. The Ni-Ti alloys used in or-thodontics are approximately 50/50 atomic percent nickel and titanium. These alloys can assume either a BCC arrangement, referred to as the austenitic phase (see Fig 20-3a), or an end-centered monoclinic arrangement, the martensitic phase (see Fig 20-3b). For the superelastic Ni-Tis, the transition from aus-tenite to martensite is induced by stress during clini-cal use. For a Ni-Ti orthodontic alloy in the austenitic phase, as the load is increased beyond point a in Fig 20-5, the crystal structure is converted to marten-site, which is referred to as stress-induced marten-site. The transition occurs at a near-constant force, producing the novel plateau region a-b in the F/∆ curve. When the load or stress is removed, the mar-tensite reverts to austenite following the unloading segment (dashed line). This mechanism is referred to as superelasticity or pseudoelasticity. The ability to undergo large elastic deections and apply low, rel-atively continuous forces is the reason for the popu-larity of austenitic Ni-Ti for the initial alignment and leveling stages of treatment.Transition temperature rangeAnother unique feature of the orthodontic Ni-Ti al-loys is that the temperature at which the austenite- martensite transition occurs is close to room and mouth temperatures. Austenitic Ni-Ti is the stable Fig 20-7 Schematic drawing of a dislocation in an atomic lattice. The dislocation (green) is an extra plane of atoms. Movement of the dislocation in response to stress is the mechanism responsible for the mechanical properties of several orthodontic alloys. 20Properties and Structures of Orthodontic Wire Materials486phase above the austenite nish temperature (Af), and martensite is stable below the martensite nish temperature (Mf). Most orthodontic Ni-Ti superelas-tic wires have an Af slightly below, and others slight-ly above, mouth temperature. Furthermore, the phase transition does not occur at one specic tem-perature but rather over a range referred to as the transition temperature range (TTR). Figure 20-8 shows the transition with temperature. At lower temperatures, the Ni-Ti is 100% martensite. As the temperature is increased, austenite starts to form at point As. With increasing temperature, the transi-tion continues until the Ni-Ti is 100% austenite at Af. With cooling, austenite starts and nishes its con-version back to martensite at the Ms and Mf tem-peratures, respectively.Importantly, the TTR can be adjusted to higher or lower temperatures with minor alterations of the Ni/Ti ratio; small substitutions of elements, such as copper for the nickel; and cold reductions and heat treatments used during manufacture of the wire. Manipulation of the TTR relative to mouth tempera-ture inuences the phases present during clinical treatment. Superelastic Ni-Ti wires respond some-what differently if the TTR is below or above mouth temperature; hence, their clinical applications are described separately.First consider Ni-Ti wires that have a TTR below mouth temperature. At room temperature, the wires are austenitic. When placed into misaligned brackets, if stress is adequate, highly stressed sec-tions of the wire convert to martensite (segment a-b in Fig 20-5). During unloading, the wire converts back to austenite following a superelastic unload-ing pattern. These wires exhibit phase transforma-tion by stress. Basically, there is no shape-memory effect or temperature-based phase transformation. Because a martensitic phase is formed by stress, heat from the mouth or food may have a small effect on the transformation back to austenite, but practically this can be ignored.Now consider superelastic Ni-Ti wires with a TTR above mouth temperature. Altering the austen-ite nish temperature (Af) between 37oC and 45oC or higher will result in varying amounts of austen-ite and martensite at average mouth temperature (37oC). Because martensite has a lower stiffness, this explains in part why the higher-TTR wires of the same composition deliver lighter forces. These wires are completely or partially martensitic at room tem-perature. When inserted into crooked teeth, they are martensite (if partly martensite, the martensite fraction can increase). Mouth temperature then transforms them completely or partially to austen-ite. Hot uids or food may also be required for full transformation to Af. In theory, one would expect complete transition at Af. Studies show that slight-ly higher temperatures may be required because of the induced stress of placing the wire into mis-aligned teeth. Advantages of superelastic Ni-Ti wires with TTRs higher than mouth temperature include the fol-lowing: For the same cross section and composition, they deliver lighter forces; the shape-memory ef-fect can be used; and the patient can control force and potential pain by drinking hot or cold uids. The disadvantage is greater variability in predicting and delivering the forces. For example, two super-Fig 20-8 The austenite-martensite Ni-Ti transitions occur over a temperature range (TTR). At lower temperatures, the structure is martensite. When the temperature is increased to As, austenite starts to form, and the transition is 100% complete at temperature Af. On cooling, martensite starts to reform at Ms, and the transition is complete at Mf. 487Materials for High Esthetics and Biomechanical Performanceelastic 0.017 × 0.025–inch Ni-Ti wires with different Af could produce as much as 1,000-gmm different bending moments when deected the same 40 de-grees. The forces produced by the same wire can also vary with the temperature experienced in the mouth. For example, the bending moment produced by a 35oC copper Ni-Ti wire deected 50 degrees can vary from 0 to 1,200 gmm with temperature increas-es from 22oC to 60oC. Besides the relative amount of phases, other mechanisms may contribute to these force variations because the Ni-Ti systems are metal-lurgically very complex.Control of the TTR around mouth temperature also can be used to facilitate placement of Ni-Ti wires. With an appropriately selected TTR, the wire can be cooled to its softer martensitic state and in-serted into the brackets; it then transitions to the higher-stiffness austenitic form as the component warms to mouth temperature.Shape memoryThe superelasticity mechanism discussed above is produced by both stress- and temperature-induced phase changes. The austenite-martensite phase transition can also be driven primarily by tempera-ture, which is the mechanism that produces the well-known shape-memory effect of Ni-Ti alloys. Through processing by the manufacturer, a par-ent shape can be produced while the alloy is in the higher-temperature, austenitic phase. Reducing the temperature causes the transition to martensite, which involves a shift of atoms to a different lat-tice structure, but the atoms retain their nearest neighbors. While in the martensitic state, the wire is formed to another shape. Upon subsequent heat-ing back into the austenite region, the atoms shift back to their original positions, returning the wire to its parent shape. The original intent of Ni-Ti al-loys in orthodontics was to use the shape-memory effect, but that application is still being perfected. Two examples of the application of shape-memory principles for orthodontic wires are the cooling of wires to make insertion easier and patient control over force by hot and cold liquids. Nevertheless, most Ni-Ti products primarily use the high spring-back associated with the superelastic mechanism of austenitic Ni-Ti and the low stiffness associated with bond stretching of martensitic Ni-Ti.Because the current orthodontic Ni-Ti wires may use superelasticity or shape memory, and because these mechanisms are dependent on temperatures and stresses that vary during clinical use, the forces produced can vary and can be difcult to accurately predict. Accordingly, research and development re-garding these concepts continue, and comparisons of products from various manufacturers are difcult.Materials for High Esthetics and Biomechanical PerformanceWhile biomechanical performance has been the pri-mary driver for the development of new materials for orthodontic therapy, there is continually grow-ing interest in improved esthetic appliances. Lin-gual orthodontics and aligners are two approaches, but each has limitations and challenges compared with traditional labial treatment. The development of clear and translucent ceramic brackets provides improved esthetics but still requires use of metal-lic archwires. Coated metal wires are available, but these are still opaque, can wear or peel during forming or clinical use, and affect frictional forces. To achieve high-quality esthetic wires, the ortho-dontic eld has begun to adopt composites and will likely employ advanced, high-strength polymers in the future. Meeting the challenge of high esthetics and biomechanical performance in an archwire will require more sophistication by the profession and an understanding of these two classes of materials.Composites in orthodontic wiresIn orthodontics, one approach for achieving both esthetics and performance might be composites. Composite materials combine the favorable features of two different constituents to obtain a preferred combination of properties. For example, dental re-storative composites obtain strength from the ller particles and handling and esthetics from the resin matrix. Composite orthodontic wires likely would use long, continuous bers in a polymer matrix, as shown with a combined scanning electron micrograph and schematic drawing in Fig 20-9. Fiber-reinforced com-posites (FRCs) were rst attempted in dentistry in the 1960s to 1980s to strengthen or construct den-ture bases, prosthodontic frameworks, orthodontic retainers, and splints. These applications were not broadly accepted in the profession, in part because the high theoretical mechanical properties were never achieved in practice. In the 1990s, it was rec-ognized that the methods used to incorporate bers into dental appliances were not developing the nec-essary coupling between the bers and the resin matrix and that the ber volume fraction was low. An alternative approach was suggested of rst im- 20Properties and Structures of Orthodontic Wire Materials488pregnating the ber bundles with the resin to en-sure coupling, high ber volume, and uniform ber distribution, then subsequently forming this pre-preg (an uncured or partially cured wire) into the desired shape. This method was successful, and FRCs are now used in orthodontics for splints, endodontic posts, and anchorage components (Fig 20-10). This brief history is instructive, because the profession is likely to use various methods of fabricating compos-ites in developing different approaches for their use in active orthodontic appliances.Even with an understanding of composite tech-nology, there has been very limited use of FRCs in archwires because the current materials are brit-tle. The same features of good coupling and high-strength continuous glass bers that impart high strength and rigidity also limit the amount the wires can be deected, and FRC wires with small cross sections used during the initial stages of treatment have been prone to failure. Nevertheless, because of their transparency or translucency, excellent esthet-ics can be achieved, and FRC orthodontic wires and systems are commercially available.Flexure properties of FRC orthodontic wires have been reported, although there are only a few pub-lished studies.2 Force levels are broadly in the low end of the range typically used in orthodontics and similar to those of Ni-Ti alloys, although comparisons need to be done with care because superelastic Ni-Ti wires are nonlinear during unloading (see Fig 20-6). The modulus of orthodontic composite wires has been re-ported in the range of up to 30 to 40 GPa.2 Even with clinically relevant force levels and excellent esthet-ics, the use of composites is challenging because of brittleness, lack of clinical formability, and anisotropy associated with unidirectional continuous bers. An alternative could be use of very high–strength poly-mers that may not require ber reinforcement. Polymers in orthodontic wiresLooking forward into the not-too-distant future, advanced, high-strength polymers may be able to provide esthetics, good biomechanical perfor-mance, and formability in an orthodontic archwire. A polymer without traditional glass or ceramic ber reinforcement could be transparent or translucent and easily pigmented if a tooth color is preferred. The fundamental concern with polymers has been insufcient strength or rigidity to deliver the neces-sary forces for active tooth movement. Thermoplas-tic polymers would likely be used in this application because of their potential for formability.The mechanical properties of thermoplastics are dependent on the chemistry, architecture, and ri-gidity of the polymer chains. One method for en-hancing chain rigidity is to add phenylene groups or rings, as represented in Fig 20-11. A polymer with all phenylene groups in the backbone would have very good rigidity and strength but could not be pro-cessed. However, by proper addition of exible seg-ments, such as those indicated by [A]m in the gure, polymers with exceptional mechanical properties and processability can be developed. In a manner analogous to ber composites, such polymer archi-tectures could be considered as molecular reinforce-ment or self-reinforcement, because the rigid ring segments provide the increased properties.The chemical structure shown in Fig 20-11 is ge-nerically polyphenylene. Recent work with poly-Fig 20-9 Schematic drawing of a unidi-rectional, continuous ber-reinforced com-posite. The end of the rectangular wire is represented by a scanning electron micro-graph that shows the individual glass bers held in position by the surrounding resin matrix. Properties vary with ber content.a bFig 20-10 (a and b) Examples of FRCs used as anchorage. 489Materials for High Esthetics and Biomechanical Performancephenylene polymers has shown that wires can be ex-truded with cross-sectional dimensions appropriate for orthodontic therapy. The modulus and ultimate strength of the wires was approximately 5.3 GPa and 160 MPa, respectively.3 Importantly, in exure the stiffness of a 0.021 × 0.030–inch polyphenylene wire was 190 gmm/degree, and the maximum moment was 1,710 gmm, both in the range of current small- diameter β-Ti and Ni-Ti wires. As an example, Fig 20-12 shows that a 0.021 × 0.030–inch poly-phenylene wire has about 80% of the stiffness and maximum moment of a 0.016-inch β-Ti wire.Besides attention to adequate stiffness and force levels, the use of nonreinforced polymers as ortho-dontic archwires would require characterization of stress relaxation. This property measures any de-crease in force over time. None of the metal alloys and reinforced composites currently used in ortho-dontic therapy exhibit any clinically signicant de-crease in force after activation simply due to time. The forces from active wires decrease due to tooth movement. Figure 20-13 schematically illustrates stress relaxation. The red line shows the response for a metal. After activation, there is no decrease in force simply due to time; the wire is stable. How-ever, a polymer may experience time-dependent re-duction in force, as represented by the blue line.Polymer chains can respond to stress with three molecular mechanisms—elastic, viscoelastic, and plastic deformation—as schematically illustrated in Fig 20-14. As with other materials, simple bond stretching and recovery provides typical elastic be-havior (Fig 20-14a). However, under stress and over time, the long polymer chains may change their con-Fig 20-11 Diagram of a generic polyphenylene polymer. The ring structure increases mechanical properties, and the exi-ble segments ([A]m) improve processing and ductility.Fig 20-12 Moment-deection curves show that the exure properties of a 0.021 × 0.030–inch polyphenylene wire are about 80% of those of a 0.016-inch β-Ti wire.Fig 20-13 Use of a polymeric wire in orthodontics would require consideration of stress relaxation. After activation, metal wires, represented by the red line, do not experience any clinically signicant decrease in force simply due to time. Polymeric wires would likely experience some decay, as sug-gested by the blue line.Fig 20-14 Schematic diagrams of the behavior of polymer chains re-sponsible for different aspects of viscoelasticity. (a) Simple bond stretch-ing accounts for the elastic behavior. (b) Chain uncoiling, extension, and reversible coiling produces a reversible, time-dependent response. (c) Irreversible sliding of chains past each other produces permanent defor-mation.cbaDeection (degrees)Moment (gmm)0.016-inch β-Ti0.021 × 0.030–inch polyphenyleneForceTime0ActivationMetalPolymeric 20Properties and Structures of Orthodontic Wire Materials490formation, uncoil, and extend. If this extension re-verses itself when the load is removed, then with time any deection is recovered (Fig 20-14b). This is time-dependent, viscoelastic behavior. If the poly-mer chains irreversibly slide past each other, the polymer experiences permanent, plastic deforma-tion (Fig 20-14c). Of course, a benet of the ability to plastically deform is the potential formability of a polymeric wire, as shown by the complex bends in Fig 20-15.Further work will be necessary before polymeric wires will be commercially available for orthodontic therapy. Additional relevant properties will need to be evaluated and techniques optimized to take ad-vantage of the manufacturing and handling charac-teristics that will be novel in orthodontics. However, given the historical trend of ever-improving mate-rials, it seems that polymeric wires are likely in the future for the profession.References1. Burstone CJ, Qin B, Morton JY. Chinese NiTi wire—A new orthodontic alloy. Am J Orthod 1985;87:445–452. 2. Goldberg AJ, Burstone CJ. The use of continuous ber re-inforcement in dentistry. Dent Mater 1992;8:197–202. 3. Burstone CJ, Liebler SA, Goldberg AJ. Polyphenylene poly-mers as esthetic orthodontic archwires. Am J Orthod Den-tofacial Orthop 2011;139(4 suppl):e391–e398.Recommended ReadingBurstone CJ, Goldberg AJ. Beta titanium: A new orthodontic alloy. Am J Orthod 1980;77:121–132.Iijima M, Ohno H, Kawashima I, Endo K, Mizoguchi I. Me-chanical behavior at different temperatures and stresses for super elastic nickel-titanium orthodontic wires having differ-ent transformation temperatures. Dent Mater 2002;18:88–93.Krishnan M, Singh JB. A novel B19’ martensite in nickel titani-um shape memory alloys. Acta Mater 2000;48:1325–1344.Sakima MT, Dalstra M, Melsen B. How does temperature inu-ence the properties of rectangular nickel-titanium wires? Eur J Orthod 2006;28:282–291.Santoro M, Nicolay OF, Cangialosi TJ. Pseudoelasticity and thermoelasticity of nickel-titanium alloys: A clinically oriented review. Part I: Temperature transitional ranges. Am J Orthod Dentofacial Orthop 2001;119:587–593.Yoneyama T, Doi H, Hamanka H, Okamoto Y, Mogi M, Miura F. Super-elasticity and thermal behavior of Ni-Ti alloy orthodon-tic arch wires. Dent Mater J 1992;11:1–10.Zhang XY, Sehitoglu H. Crystallography of the B2 ➔ R ➔ B19’ phase transformations in NiTi. Mater Sci Eng 2004;374:292–302.Fig 20-15 (a to c) The ability of a polymeric wire to undergo large amounts of plastic deformation allows forming of complex bends.a b c

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