Ashok Sir Arch Wire

8/9/2019 Ashok Sir Arch Wire

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I ntroduction

Wire is one of the important components of all most all orthodontic

appliances. Practically all orthodontic forces for which appliances are used exert

forces by means of wires – not just any wire, but a wire properly selected in size,

shape, material, properties and properly bent to exert the desired force. An

understanding of the well balanced relationship that exists between the applied

techniques and the basic principles, leads to a broader application of skills to serve the

need of orthodontics.

An orthodontist spends much of his professional career handling wires and the

success or failure of many forms of treatment depends upon the correct selection of

wires, possessing adequate properties combined with careful manipulation beside

bracket and auxillaries. The search for correct materials has continued from the

beginning of dental art to the present time. Through the ages, dentistry has been

dependent to a great degree on the advances made by the contemporary art and

sciences for improvements in materials.


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The materials used by orthodontists have changed rapidly in recent years and

will continue to do so in the future. As esthetic composite arch wires are introduced,

metallic arch wires are likely to be replaced for most orthodontic applications in the

same way as metals have been replaced by composites in aerospace industry.

Arch wires are reviewed in the order of their development, with emphasis on

specific properties and characteristics, such as strength, stiffness, range, formability

and weldability. Because an ideal material has not yet been found, arch wires should

be selected within the context of their intended use during treatment.

Over the last century, material science has made rapid progress. This has been

evident in our day to day life also. Orthodontics, particularly, has benefited largely

from this. In this branch of dentistry, not only have the materials been improved, but

also the philosophies have changed. Orthodontics has come a long way since the days

of the E-arch and various removable appliances used in the early 20th century. With

the introduction of the Edgewise appliance, newer materials have introduced in order

to make the most of these appliances. Wires which had good formability, increased

resilience and low cost were obviously favoured. This was probably the reason why

stainless steel (and Elgiloy) prevailed over the noble metal alloys.

The need of the Begg appliance was quite different from that of the traditional

edgewise appliance. This led Begg and Wilcock to produce a variety of stainless steel

that would provide low continuous forces over a long period of time. The Nickel-

Titanium(Ni-Ti) alloys introduced in the 1970’s showed some remarkable properties

of superelasticity and shape memory, although these could not be exploited clinically

at that time. The wires had limited formability, but could still be used in the


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traditional edgewise appliance. The next generation of NiTi wires benefited a lot by

the pre adjusted edgewise appliance.

This appliance required lesser amount of bends incorporated into the wire, and

the A- NiTi’s perfectly suited this. However, introduction of the TMA wires bridged

the gap between stainless steel and Nickel Titanium alloys wires, with properties that

were intermediate to the two of these alloys.

Thus, one can see how the appliance philosophies and material science

progress is closely interrelated. All these wire alloys that were introduced and the

newer ones have individualistic and unique properties associated with them. In order

to use the newer wires, it is important to know as to why they have specific properties.


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Review of Literature

Rapid strides have been made in the field of arch wire materials. The urge for

better performance has resulted in the development of newer orthodontic wires with

promising physical properties.

PIERRE FAUCHARD, the father of modern dentistry in 1723 developed

what is probably the first orthodontic appliance in evolution of fixed orthodontic

appliance. It was called as Bandolet or Bow. It was flat piece of metal scalloped out

for the ideal position of the teeth. The teeth were ligated towards their positions. This

appliance was very heavy and unwieldy. It was also designed to expand the arch,

particularly the anterior teeth. FAUCHARD said “If the teeth are much out of line and

cannot be corrected by means of thread, it is necessary to use a band of silver or gold.

The width of band should be less than the height of the teeth to which it is applied.

The band should neither be too stiff or too flexible. Two holes are made at each end,

and a thread passing partially forms a loop and by the pressure and support given the

inclined teeth will be made upright”.2


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In 1757, Etienne Bourdet (1722 to 1789), the dentist to the king of France,

advocated the Fauchard method but went a step further by recommending only gold

strips on the labial surface of upper arch and lingual surface of lower arch. He wrote in

his book that “The strings should be removed and retightened twice a week, until the

teeth have resumed their proper position – that is to say, until the teeth of upper Jaw are

drawn forward so that no part of them is hidden behind those of the lower jaw”.2

Leonard Koecker (1728 to 1850) in 1826, practicing in Philadelphia,

advertised that “He supplies ligatures to teeth of an irregular position”.2

Samuel S. Fitch MD, whose book entitled A system of dental surgery,

published in 1829, devoted a significant amount of information on irregularities of the

teeth. He was also the first one to classify malocclusion. His treatment consists of

“Application of an instrument adapted to arch of the mouth, fastening a ligature on the

irregular tooth and removing the resistance of the lower teeth by placing some

intervening substances between the teeth of upper and lower jaw, so as to prevent

them from completely closing”.2

Shearjashub Spooner (1809 to 1859) in 1838 found various types of

treatments, such as use of gold and silver plates to exert a gentle and continued

pressure to correct irregularities of teeth.2

William Lintott in 1841 described a bite opening appliance, which consisted

of a labial arch of a light bar of gold or silver passed around front surfaces of teeth by

means of ligatures (known as Indian twist) and the necks of irregular teeth with

pressure applied for movement.2


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As early as in 1871 William E. Magill (1825 to 1896), was first to use

cemented bands on the teeth by oxychloride of zinc cement. It was on the foundation

of this cemented tooth band and circumferential arch wires that modern orthodontic

appliance have developed.2

In 1887, Dr. Angle introduced the round labial arch wire which was

supported by clamp bands on molar teeth. It also was an expansion arch and teeth

were ligated towards their preplanned arch. If molar expansion was desired the arch

wire was expanded. The appliance is commonly referred to as E (expansion) arch.

As demand increased for more and better control of the teeth, bands were added to

anterior teeth with vertical tubes placed over them. Like this the pin and tube

appliance was developed.

In 1916 with the advent of ribbon arch, the E arch gave way to flat wire 0.022″

x 0.036″ placed against the teeth. This flat flexible wire was molded to fit the

malocclusion and was held in close approximation to the teeth by a bracket that opened

occlusally. It has excellent rotating ability but lacked the power to tip the teeth.

In 1908, Dr. P.R. Begg designed an appliance for moving roots of teeth.

In 1929 Dr. Angle introduced an appliance that engages the teeth edge wise

by way of new bracket that opened bucally and used flat wires of 0.028″ dimension.

Thus the edgewise appliance was introduced.

It could be observed that in Angle’s orthodontic appliance, the arch wires in

each succeeding mechanism was thinner than the immediately preceding mechanism;

so that the amount of forces delivered for tooth movement became less in each later


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mechanism. This indicates that Angle was aware that the tooth moving forces

delivered by his earlier forms of orthodontic appliance were too great. This reduction

of tooth moving forces in each new orthodontic mechanism permitted greater control

of tooth movement. It made possible to move the teeth rapidly and reduced the pain

that patient had to bear during treatment.

Up to 1930,s the only orthodontic wire available was made of gold. In 1929

Lucien de Costa a Belgian and editor of Archives of orthodontics introduced

austenitic stainless steel orthodontic wire with greater strength, high modulus of

elasticity, good resistance to corrosion and low cost.

It was in between 1903 and 1921 that Harry Brearley of Sheffield , F.M.

Becket of USA, Beune Strauss and Edward Maurer of Germany shared the honor

for the development of the materials.

In 1937, Atkinson introduced Atkinson, s universal appliance. He used two

different forms of labial wire, one rectangular and one round and was designed to bring

about every tooth movement possible. A significant advancement in orthodontic

materials was made in late 30,s and 40,s when stainless steel wires became widely

available. The cobalt alloys were simultaneously developed in the mid century and this

has physical properties very similar to that of stainless steel. They had an advantage that

they could be supplied in softer and more formable state and then could be hardened by

heat treatment. The procedure increases its strength significantly.

In 1952 Dr. Begg in collaboration with Mr. A.J.Willcock sought to develop

tensile wire materials that were thin enough to distribute forces at an optimum level


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for tooth movement over a considerable period of time, over a long distance and with

minimal loss of force intensity. The wire was thick enough to resist masticatory stress.

The diameter of wire initially produced was progressively decreased from the thicker

diameter to 0.018″ to 0.014″ arch wire.4

Then came the most talked Niti wire which was invented in 60,s by William

F. Buchler , a research metallurgist at the Naval Ordinance Laboratory in Silver

Spring, Maryland (now called as Naval Surface Weapons Center). He did extensive

research and published his findings on the properties and uses of his new alloy. The

name Nitinol is an acronym derived from elements which comprises the alloy, Ni

from nickel, Ti from titanium and nol from Naval Ordinance Laboratory.

Niti was introduced to orthodontics by Andreasen and his associates. They

were attracted to unique properties of Niti alloy, such as high elastic limit and low

modules of elasticity.

In 1971, they reported the results of their investigation for clinical use and

subsequently Unitek Corporation started producing this wire for clinical use under

the trade name of Nitinol. It has an excellent spring back property but does not

possess shape memory or super elasticity because it has been manufactured by a work

hardening process.

Later developments related to Niti alloy came from China in Beijing in

General research institute for Non-ferrous metal in 1978, by DR. Hau-Chang Tien

and his colleagues with Niti a new super elastic orthodontic wire with high spring

back and low stiffness properties.9


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In the same year Furukawa electric company Ltd of Japan produced a new

type of Japanese Niti alloy possessing properties of excellent spring back, shape

memory and super elasticity.29

In 1980, Dr. Andreasen tested thermodynamic nitinol wires and introduced

them to clinical orthodontics. These wires can return to previously set shape when

heated to their transition temperature range (TTR). He was the first person to suggest

the use of shape changes in Nitinol wires to apply forces to the teeth in order to move

them orthodontically.

At around the same time in 1980, Charles J. Burstone and A. Jon Goldberg,

introduced new Beta-titanium alloy (Titanium-molybdenum alloy) in clinical use of

orthodontics. It has a unique balance of low stiffness, high spring back, good

formability and weldability which indicates its use in a wide range of clinical

applications.8

In 1985, Dr. C.J. Burstone reported the development of Chinese Niti alloy

and in 1986 Miura Fetal reported Japanese Niti alloy. These two alloys have a basic

austenitic grain structure and have the advantage of a transition in the internal

structure without requiring a significant temperature change to do this.

In 1988 Mr. A.J. Willcock Jr. of Australia developed a much harder, near

alpha-phase titanium alloy comprising of 6% Aluminum and 4% Vanadium for

orthodontic purposes.4 He also started the production of ultra high tensile stainless steel

fine round wire, supreme grade as per the request of Dr.Mollenhauer of Melbourne.

The wire was initially in the0.010″ diameter and was further reduced to 0.009″.


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In 1990 John J. Hudgins, Michael D. Bagby and Leslie C. Erickson studied

the effect of long term deflection on permanent deformation of Nickel- Titanium.17

In 1991 Sunil Kapila, Gary D. Richhold and Etal investigated the Nickel

titanium alloy to determine the effect of clinical recycling on load deflection

characteristics and surface topography of Nickel-titanium alloy.

In 1992 Glen A. Smith, J.A. Von Fraunhofer , Glenn R.Casey studied the

effect of clinical use and various sterilization procedures on three types of Nickel-

titanium and one type of Beta-titanium and stainless steel arch wire. The various

procedure included disinfection alone and in conjugation with steam autoclave, dry

heat and cold solution sterilization.26

In 1992, the same year, OPTIFLEX an aesthetic arch wire, was introduced to

orthodontics by Tallas. It is made up of clean optical fiber and has unique mechanical

properties.36

In 1995 Charles J. Burstone demonstrated Titanium molybdenum alloy

(TMA) with ion implantation. A low coefficient of friction is usually desirable in

orthodontic arch wire. Studies have shown that Nickel titanium and TMA have higher

coefficient of friction than stainless steel.

In case of TMA, the friction is probably high due to its relative softness

compared to harder stainless steel bracket. Ion implantation increases its hardness and

reduces coefficient of friction of TMA wire.8


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In 1995, the same year Rohit Sachdeva and Suchio Miyasaki introduced

copper-Niti alloy in family of Niti. It’s an alloy of copper, nickel, titanium and

chromium.

Recently in 2001, Dead Soft Security Arch wires has been introduced by

Binder and Scott . These arches are bend to lie passively in all attachments.5


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Classification

Arch wires can be broadly classified according to chemical composition,

microstructure and mechanical properties.

1) According to Materials used

GOLD ARCHWIRES

STAINLESS STEEL ARCHWIRES

AUSTRALIAN ARCHWIRES

CHROME COBALT NICKEL ALLOY ARCHWIRES

JAPANESE NITI ARCHWIRES

CHINESE NITI ARCHWIRES

ALPHA-TITANIUM ALLOY ARCHWIRES

COPPER-NITI ALLOY ARCHWIRES

NICKEL SILVER ALLOY ARCHWIRES

FORSTADENT TITANOL ARCHWIRES


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OPTIFLEX ARCHWIRES

DEAD SOFT SECURITY ARCHWIRES

NICKEL TITANIUM ARCHWIRES

CONVENTIONAL

PSEUDOELASTIC

THERMODYNAMIC

2) According to Cross- section

ROUND

RECTANGULAR

ROUNDED RECTANGULAR

SQUARE

BRAIDED

MULTISTRANDED

3) According to Diameter

0.008″ to 0.045″ FOR INTRA ORAL APPLIANCES

0.045″ to 0.60″ FOR EXTRA ORAL APPLIANCES


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per unit area, whereas strain is the internal distortion produced by the load, defined as

deflection per unit area.

Orthodontic arch wires and springs

can be considered as beams, supported

either only on one end (e.g. a spring

projecting from a removable appliance) or

from both ends (a segment of an arch wire

spanning between attachments on adjacent

teeth). If a force is applied to such a beam,

its response can be measured as the deflection produced by the force. Force and

deflection are external measurements.

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CANTILEVER- A

UPPORTED BEAMS-B


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STRESS STRAIN DIAGRAM

In tension, internal stress and strain can be calculated from force and

deflection by considering the area and length of the beam. For Orthodontic purposes,

three major properties of beam materials are critical in defining their clinical

usefulness i.e. strength, stiffness and range. Each can be defined by appropriate

reference to a force deflection or stress strain diagram.

Three different points on a stress-Strain diagram can be taken as representative

of the strength of a material. Each represents, in a somewhat different way, the

maximal load that the material can resist. The most conservative measurement is the

proportional limit, the point at which any permanent deformation is first observed. A

more practical indication is the point at which a deformation of 0.1% is measured; this

is defined as the yield strength. The maximum load that the wire can sustain- the

ultimate tensile strength is reached after some permanent deformation and is greater

than the yield strength. Since this ultimate strength determines the maximum force the

wire can deliver if used as a spring, it is important clinically, especially since yield

strength and ultimate strength differ much for titanium alloys. Strength is measured in

stress units (gm/cm square)

Stiffness and springiness are reciprocal properties.

Springiness = 1/stiffness


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FORCE DEFLECTION CURVE

STRESS STRAIN DIAGRAM

Each is proportional to the slope of the elastic portion of force deflection

curve. The more horizontal the slope, the springier the wire; the more vertical the

slope, the stiffer the wire.

Range is defined as the distance that the wire will bend elastically before

permanent deformation occurs. This distance is measured in mm. If the wire is

deflected beyond its yield strength, it will not return to its original shape, but

clinically useful spring back will occur unless the failure point is reached. This spring

back is measured along the horizontal axis as shown in figure.

In many clinical situations, orthodontic wires are deformed beyond their

elastic limit. Their spring back properties in the portion of the load-deflection curve


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are between the elastic limit and the ultimate strength, therefore, are important in

determining clinical performance.

These three major properties have an important relationship.

Strength = Stiffness X Range.

Two other characteristics of some clinical importance can also be illustrated

with a stress strain diagram; resilience and formability. Resilience is the area under

the stress- strain curve out to the proportional limit. It represents the energy stored

capacity of the wire, which is a combination of strength and springiness. Formability

is the amount of permanent deformation that a wire can withstand before failing. It

represents the amount of permanent bending the wire will tolerate before it breaks.

The properties of an ideal wire material from orthodontic purposes can be

described largely in terms of these criteria:

High Strength

Low stiffness

High range

High formability.

In addition, the material should be weldable or solderable so that hooks or

stops can be attached to the wire. It should also be reasonable in cost. In

contemporary practice , no one arch wire material meets all these requirements , and

the best results are obtained by using specific arch wire materials for specific

purposes.


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WIRE CHARACTERSTICS OF CLINICAL RELEVANCE

Several characteristics of orthodontic wires are considered desirable for

optimum performance during treatment. These include a large spring back, low

stiffness, high formability, high stored energy, biocompatibility, environment

stability, low surface friction and the capability to be welded or soldered to

auxiliaries and attachments. A brief description of each of these desirable wire

characteristics is provided.

1) SPRING BACK

This is also referred to as maximum elastic deflection, maximum flexibility

and range of activation or working range.

Spring back is related to the ratio of yield strength to the modules of elasticity

of the material. (Ys/E). Higher spring back values provide the ability to apply large

activation with a resultant increase in working time of the appliance.

This in turn implies that fewer arch wire changes or adjustments will be

required. Spring back is also a measure of how far a wire can be deflected without

causing permanent deformation or exceeding the limits of the material.

2) STIFFNESS OR LOAD DEFLECTION RATE

This is the force magnitude delivered by an appliance and is proportional to

the modulus of elasticity. Low stiffness provides the ability to apply lower forces, a

more constant force overtime as the appliance experiences deactivation and greater

ease and accuracy in applying a given force.


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3) FORMABILITY

High formability provides the ability to bend a wire into desired

configurations such as loops, coils and stops without fracturing the wire.

4) MODULUS OF RESILIENCE OR STORED ENERGY

This property represents the work available to move the teeth. It is reflected by

the area under the line describing elastic deformation of the wire.

5) BIOCOMPATIBILITY AND ENVIRONMENTAL STABILITY

Biocompatibility includes resistance to corrosion and tissue tolerance to

elements in the wire. Environmental stability ensures the maintenance of desirable

properties of the wire for extended periods of time after manufacture. This in turn

ensures a predictable behavior of the wire when in use.

6) JOINABILITY

The ability to attach auxiliaries to orthodontic wires by welding or soldering

provides an additional advantage when incorporating modifications to the appliance.

7) FRICTION

Space closure and canine retraction in continuous arch wire techniques involve

a relative motion of bracket over wire. Excessive amount of bracket / wire friction

may result in loss of anchorage or binding accompanied by little or no tooth

movement. The preferred wire material for moving a tooth relative to the wire would

be one that produces the least amount of friction at the bracket / wire interface.


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M anufacturing

All stainless steel orthodontic wires are produced with the help of standard

formulas based on specifications of the American Iron and steel Institute.

The physical properties of metals are influenced at every step in production,

beginning with the selection and melting of alloying metals.

INGOT

Dentists are so used to forget that an orthodontic wire is actually a modified

cast. One of the critical steps in wire making is pouring the molten alloy into a mold

to produce an Ingot.

This Ingot is far from being a uniform chunk of metal. Like any casting it will

have varying degree of porosity and inclusions of slag in different part.

A magnified view of inside of Ingot would show it, to be made up of crystals

of component metals. In metallurgical terminology these crystals are usually called


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grains, and it is this granular structure which controls many of the mechanical

properties.

Grains in a crystal are found in definite patterns typical of individual metals,

but they are far from perfect because of conditions under which they must form.

When the Ingot is cooling and solidifying, many different grains are forming at once.

These growing crystals crowd and surround one another, so that the ingot

becomes a mesh work of many irregularly shaped grains of different materials. The

size and distribution of these grains are very dependent on the rate of cooling and the

size of the ingot.

The cooling and pouring processes affect the porosity as well as grain structure.

Porosity in the ingot comes from either of two sources, gases that are either dissolved in

the metal or produced by chemical reactions within the molten mass from bubbles

which are trapped in metal. As the ingot cools and shrinks, the late cooling interior

section shrinks inside an already hardened shell. This shell does not permit the volume

to adjust enough to the shrinkage, so additional voids of the vacuum results. So, before

further processing begins the ingot is trimmed to remove the undesirable parts.

The microstructure of a metal is the very basic of its physical properties and

mechanical performance and every step in production is directed at getting the most

out of the original grain structure of the ingot.

ROLLING

The first mechanical step in processing is rolling the ingot into a long bar. This

is done by a series of rollers which gradually reduce the ingot to a relatively smaller


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diameter. Through all this rolling and later processing into the final wire, the different

parts of original ingot never lose their identity.

The metal that was on the outside of the ingot forms the finest wire. Wire is

actually a grossly distorted ingot, thus it is easy to see that different pieces of wires

from the same batch can differ depending upon which part of ingot they came from.

The individual grains of the ingot also keep their identity through the rolling

process until certain heat treatment is applied. Each grain is elongated in the same

proportion as the Ingot. The squeezing, massaging action of rolling the Ingot has a very

important effect on the grain structure, actually increasing the strength of the metal.

Where the original crystal fitted together rather indifferently with gaps and

voids scattered among them, the mechanical action of rolling, forces them into long,

finger like shapes that are closely meshed together. This causes an increase in the

hardness or brittleness of the metal, as the grains are forced to interlock even more

highly with one another. This is a form of work hardening. Even the atoms which

make up the crystal structure are forced into new positions, filling in gaps and

irregularities that may have been left in original crystals.

Each pass through the rollers, increases this work hardening and finally the

structure becomes so locked up that it can no longer adjust enough to adapt to the

squeezing of the rollers. If rolling is continued beyond this point the surface will start

to show many small cracks and begin to crumble. Before this happens the rolling

process is stopped and the metal is annealed by heating to a suitable high temperature.

At annealing temperature the atoms become mobile enough to move about within the


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mass, breaking up the tight crystalline structure. When the metal is cooled again, the

annealed structure resembles that of the original casting but in more uniform form.

Grains size can be controlled in annealing by adjustment of the time and temperature

of annealing and rate of cooling.

DRAWING

After the ingot has been reduced to a fairly small diameter by rolling, it is

reduced to its final size by drawing. This a more precise process in which the wire is

pulled through a small hole in a die. This hole is slightly smaller than the original

diameter of the wire so that the walls of the die squeeze the wire uniformly from all

sides, as it passes through. This reduces the wire to the diameter of the die. Drawing

the wire subjects the entire surface of the wire to the same pressure instead of

squeezing from only two sides as in rolling.

Drawing is much precise process than rolling, but the effect on grain structure

is much the same. Before it is reduced to orthodontic wire/size, the wire must be

drawn through many series of dies and annealed several times along the way to

relieve work hardening.

These intermediate annealing is very important for strength and especially to

resistance to breakage. The purpose of heating and cooling a large coil of wire so that

all parts are treated alike is not as easy as it may seem. It must be done slowly to

prevent the outer coils from being heated more than those on the inside and

temperature must be carefully controlled.


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Even with the most careful procedures, situations can arise in which one side

of the coil or the inner or outer part will be affected differently. Variations such as

these can create many problems in sampling for quality control.

The actual no of drafts through the dies as well as frequency of annealing

depends on the alloy being drawn. Gold is extremely ductile and can be reduced

considerably with each draft. Ordinary carbon steel requires many more steps than

gold and stainless steel requires many more than carbon steel. Gold work hardness

slowly, so that it also needs less frequent annealing than the more rapidly work

hardening steel.

Hardness and spring properties of orthodontic wires depend almost entirely on

the effect of work hardening during manufacture. This means that the entire drawing

and annealing schedule must be carefully planned with the final size in mind. If the

metal is almost in need of another annealing at its final size, it will have maximum

work hardening and spring properties. If drawing is not carried out for enough time

after the last annealing, there will be too much residual softness.

Wires can be reduced through much of the range of orthodontic size without

an intermediate annealing. When wire is annealed in processing at one size and

different parts of the batch are then drawn to different final sizes, the smaller of these

wires will be subjected to more hardening. This effect is usually rather small and

because of different drawing schedules that are used, it is not consistent. Differences

in these cases make the smaller wire proportionally harder, which is desirable as long

as brittleness does not become excessive.


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I deal Orthodontic Alloy

The ideal orthodontic wire for an active member is one that gives a high

maximal elastic load and low load deflection rate. The mechanical properties that

determine these characteristics are elastic limit and modules of elasticity. The ratio

between the elastic limit and modules of elasticity (EL/E) determines the desirability

of the alloy. The higher the ratio, the better will be the spring properties of wire. The

orthodontist should look for alloys that have high EL,s and low E,s . For an alloy to be

superior in spring properties, it must possess a significantly higher ratio.12

By contrast, in the reactive member of an appliance not only is a sufficiently

high elastic limit required but a high modulus of elasticity is also desirable. Since it is

common practice to use the same size of slot or tube opening throughout the

treatment, it is possible to use different alloys combined in the same appliance so that

the needs of both the active and reactive members can be served.


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Four other properties of wire should be mentioned in evaluating an

orthodontic wire.

1) The alloy must have a reasonable resistance to corrosion caused by the fluids

of the mouth.

2) It should have sufficient ductility so that it will not fracture under accidental

loading in the mouth or during fabrication of an appliance.

3) It is desirable to have a wire that can be fabricated in a soft state and later heat

treated to a hard temper.

4) A desirable alloy is one to which attachments can easily be soldered.

A thorough knowledge of the mechanical and physical properties of an alloy is

important in the design of an orthodontic appliance.

WIRE CROSS SECTION TYPE

(ROUND, FLAT, SQUARE, RECTANGULAR)

A most critical factor in the design of an orthodontic appliance/wire is the

cross – section of the wire to be used. Small changes in cross-section can dramatically

influence both the maximal elastic load and the load deflection rate. The maximal

elastic load varies directly as the third power of the diameter of round wire, and the

load deflection rate varies directly as the fourth power of the diameter. It may seem

that the most obvious method of reducing the load deflection rate of an active member

is to cut down the size of the wire. The problem in reducing the size of cross-section

is that the maximal elastic load is also reduced at an high rate (d3). In the design of the

active member it is good policy to use as small as cross – section as possible consistent


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with a safety factor, so that undue permanent deformation will not occur. Beyond this,

any attempt to reduce the size of cross- section to improve spring properties may well

lead to undesirable permanent deformation.

The fact that the load deflection rate varies as the fourth power of the diameter

in round wire suggest the critical nature of selection of proper cross-section. A piece

of 0.018″ wire is not interchangeable with 0.020″ wire, for with a similar activation,

the 0.20″ wire will deliver almost twice as much force. In the selection of proper

cross-section for the rigid reactive members of an appliance, load deflection rate

rather than maximal elastic load is the prime consideration. Under normal

circumstances it is necessary to select a large enough wire cross- section, beyond the

needed maximal elastic load to have sufficient rigidity, so that a sufficiently high load

deflection rate exists.

Factors influencing load deflection

DESIGN FACTORLOAD

DEFILATION RATEMAXIMUMINCREASE

MAXIMUMDEFLECTION

Activation of wirewithout changing length

decreased No change Increase

Activation in direction oforiginal bending

- Increase Increase

Alteration of cross

section to rectangularform

If rate is maintainedas constant

Increase as 1/h Increase as 1/h

MECHANICAL

PROPERTIES OF

WIRE

MODULUS OF

ELASTICITY

PROPORTIONAL

LIMITSP/E

Cross section(round) L d 1/d

Cross section(rectangle) h h 1/h

Length/cantilever 1/L 1/L L


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A): Optimal Cross section for flexible member

Generally for multi directional activations in which the structural axis is bent

in more than one plane, a circular cross-section is the choice. The mechanical

properties of the round wire and cross-section tolerances are far superior to those of

other cross-sections. One of the problems of round wire is that, unless it is properly

oriented, activations may not rotate in the intended plane. Moreover, round wire may

rotate in the bracket and if certain loops are incorporated in wire, these can roll into

either the gingival or the check.

In cases of unidirectional activations, flat wire is the cross-section of choice as

more energy can be absorbed into a spring made of flat wire than of any other cross-

section. Flat or ribbon wire can deliver lower load-deflection rates without permanent

deformation than can any other type of cross-section. Another advantage of flat wire

is that the problems of orientation of the wire can be more simply solved than with a

round cross-section.

Flat wire can be definitely anchored into a tube or a bracket so that it will not

spin during the deactivation of given spring. Flat wire can also be used in certain

situations when considerable tooth movement is required in one plane, while limited

tooth movement in other plane.

B): Optimal Cross section for reactive member

With respect to reactive member, a square or rectangular wire would appear

superior to a round one because of the ease of orientation and greater multi directional

rigidity. This leads to more definite control of anchorage units also.


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SELECTION OF PROPER WIRE

(CROSS SECTION SIZE & ALLOY USED)

The selection of proper wire is based primarily on the load deflection rate

required in the appliance. Secondarily, it is dependent on the magnitude of the forces

& moments required. Sometimes 2 other factors can be used in selecting wire cross

section size.

1) It may be believed that increasingly heavier wires are needed in a replacement

technique to eliminate the play in a first order direction between wire and the

bracket. In an edgewise appliance, the ligature wire minimizes a great amount

of play in a first order direction, since it can fully seat in the brackets.

Therefore the clinician does not select a 0.18″ wire over 0.016″ wire primarily

because of the difference in play.

2) A wire may also be selected because it is believed that the smaller the wire the

greater will be the amount of maximum elastic deflection possible; in other

words the smaller the wire the greater it will get deflected without permanent

deformations, but maximum elastic deflection varies inversely with the

diameter of wire.

The major reason why the orthodontist should select a particular wire size is

the stiffness of the wire or its load deflection rate. In a replacement technique, for

instance, the orthodontist might begin with a 0.014″ wire that deflected over 2 mm

would give the desired force. After the tooth had moved 1 mm, the wire could be

replaced with a 0.018″ which would give almost the same force with 1 mm of

activation.


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Small differences in cross-section produces large changes in load deflection

rates, since in round wires load deflection rate varies as the fourth power of diameter.

Clinicians are interested in the relative stiffness of the wire that they use, but they

have neither the time nor the inclination to use engineering formulas to determine

their stiffness.

Therefore a simple numbering stuff has been developed, based on engineering

theory that gives the relative stiffness of wires of different cross-sections if the

material composition of wire is the same.

The cross-sectional stiffness no (Cs) uses .1 mm (0.004″) round wire as a base

of a 0.006″ wire has a Cs of 5, which means that for the same activation five times as

much form is delivered. Manufacturing variations in wires or mislabeling of wires

obviously can significantly alter the actual Cs number.

CROSS SECTIONAL STIFFNESS NUMBER OF ROUND WIRE

Cross section Cs

(m) (mm)

0.004 0.102 1.00

0.010 0.254 39.06

0.014 0.356 150.060.016 0.406 256.00

0.018 0.457 410.06

0.020 0.508 625.00

0.022 0.559 915.06

0.030 0.762 3164.06

0.036 0.994 6561.00


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CROSS SECTIONAL STIFFNESS NUMBER OF RECTANGULAR

AND SQUARE WIRE

Shape

Cross section CS

M mmFIRST

ORDER

SECOND

ORDER

RECTANGULAR 0.010 X 0.020 0.254X0.508 130.52 132.63

RECTANGULAR 0.016X0.022 0.406X0.550 1129.79 297.57

RECTANGULAR 0.018X0.025 0.457X0.035 1805.10 966.87

RECTANGULAR 0.021X0.025 0.535X0.035 2173.95 1535.35

RECTANGULAR 0.0215X0.028 0.546X0.711 3129.83 1845.37

ShapeCross section

CS M mm

SQUARE 0.016X0.016 0.406X0.406 434

SQUARE 0.018X0.018 0.457X0.457 646.14

SQUARE 0.021X0.021 0.531X0.531 1289.69

Wires with a cross-section of 0.016″has a Cs of 256, which means that for an

identical activation it will deliver 256 times as much force as a 0 .004″ round wire. For

purposes of comparison both the wire configuration and the alloy are identical and

only the cross-section varies.

In the past, wire cross-section has been varied to produce different stiff nesses.

The overall stiffness of an appliance (S) is determined by two factors; one relates to

the wire itself (Ws), and one is the design of an appliance (As):

S = Ws x As

Where S = Appliance load deflection rate

Ws = The wire stiffness

As = Design stiffness factor

In general terms,

Appliance stiffness = Wire stiffness x Design stiffness


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As the appliance design is changed by increasing wire between the brackets or

adding loops, the stiffness can be reduced as the design stiffness factor changes.

However, the orthodontist is not concerned only with ways by which wire stiffness

can be altered. Wire stiffness is determined by two factors- the cross-section and

material of the wires.

Ws = Ms x Cs

Where

Ws is wire stiffness number

Ms is material stiffness number

Cs is cross sectional stiffness number.

In general terms

Wire stiffness = Material stiffness x Cross sectional stiffness

Previously, since most orthodontists used only stainless steel with identical

modulus of elasticity, only the size of the wire was varied and no concern was given

to the material property, which determines wire stiffness.

With the availability of new materials, it is now possible to maintain the same

cross-section of wire but use different materials with different stiff nesses to produce

a wide range of forces and load deflection rates required for comprehensive

orthodontics.

A numbering system can be used to compare relative stiff nesses based on the

material. The material stiffness number (Ms) is based on the modulus of elasticity of

the material.


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Since, steel is currently the most commonly used alloy in orthodontics, its Ms

Number has been arbitrarily set at 1. Typical stiffness numbers for other alloys are

given in table. Although the modulus of elasticity is considered a constant, the history

of the wire (drawing process) may have some influence on the modulus. For practical

clinical purposes, however, the material stiffness number (Ms) can be used to

determine the relative amount of force that a wire will give per unit activation.

In addition to new alloys, braided wires have been used in orthodontics. Braids

take advantage of smaller cross-sections, which have higher maximum elastic

deflections, and in process produce wires that have relatively low stiffness. The

material stiffness numbers of representative braided wires is given in table.

MATERIAL STIFFNESS NUMBER OF ORTHODONTIC ALLOYS &

BRADED STEEL

MS

ALLOYS

S.S 1.00

TMA 0.42

Nitinol 0.26

Elgiloy blue 1.19

Elgiloy blue(Heat treated) 1.22

Braids

Twist-hex 0.18-0.20

Force -9 0.14-0.16

D-rect 0.04-0.08

Respond 0.07-0.08

The load deflection rate can be changed by maintaining wire size and varying

the load deflection rate as significantly as by altering the cross-section. Using the

principle of variable cross-section orthodontics, the amount of play between the


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attachments and the wire can be varied, depending on the stiffness required. With

small low-stiffness wires, excessive play may lead to lack of control over tooth

movement.

On the other hand if the principle of variable modulus orthodontics is

employed, the clinician determines the amount of play required before selecting the

wire. In some instances more play is needed to allow freedom of movements of

brackets along the arch wire. In other situations little play is required to allow good

orientation and effective third-order movements. Once the desired amount of play has

been established, the stiffness of wire can be produced by using a material with a

proper material stiffness. In this way the play between the wire and the attachment is

not dictated by the stiffness required but is under the full control of the operator.

The variable modulus principle allows the orthodontist to use oriented

rectangular wires or square wires in light force, as well as heavy force applications

and stabilizations. A rectangular wire orients in the bracket and hence offers greater

control in delivering the desired force system. More important, when placed in the

brackets, the wire will not turn or twist to allow the forces to be dissipated in

improper directions.

WIRE LENGTH

The length of a member may influence the maximum elastic load and the load

deflection in a number of ways depending upon the configuration and loading of the

spring. The cantilever has been chosen to demonstrate the effect of length, since the

cantilever principle is widely used in orthodontic mechanisms.


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The figure shows a cantilever attached at B with vertical force applied at A.

The distance L represents the length of the cantilever measured parallel to its

structural axis.

In this type of loading the load deflection rate will very inversely as the third

power of the length; in other words, the longer the cantilever the lower the load

deflection rate. The maximal elastic load varies inversely as the length of the

cantilever. Once again, the longer the cantilever the lower the maximal elastic load.

Increasing the length of cantilever is a better way to reduce the load deflection

rate than is reducing the cross-section. Increasing the length of the cantilever

markedly reduces the load deflection rate; yet the maximal elastic load is not radically

altered, since it varies linearly with the length. Adding length within the practical

confines of the oral cavity is an excellent way of improving spring properties.

Increasing the length of a wire with vertical loops is one of the more effective means

of reducing load deflection rates for flexible members and at the same time, only

minimally altering their maximal elastic loads. However there are limitations in how

much the length can be increased. The distance between brackets in a continuous arch

is predetermined by tooth and bracket width. Vertical segments in the wire are limited

by occlusion and the extension of the muco-buccal fold.


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AMOUNT OF WIRE

Additional length of wire may be incorporated in the form of loops and helices

or some other configuration. This tends to lower the load deflection rate and increases

the range of action of the flexible member. The maximal elastic load may or may not

be affected.

When a member is designed that incorporates additional wire, it is necessary

to locate properly the parts of the configuration where additional wire should be

placed and to determine the form that the additional wire should take.

If location and formation are properly done, it should be possible to lower the

load deflection rate without altering the maximal elastic load merely by adding the

least amount of wire that will achieve these ends.

The optimal place for additional wire is at cross-sections where bending

moment is largest. In the case of cantilever the position for additional wire would be at

the point of support, since here the bending moment is the greatest, almost 1000 gm.

Helical coils can be used to reduce the load deflection rate. The figure

illustrates the proper positioning of helical coil for this purpose. The load deflection

rate is maximally lowered for the given amount of wire used if the helix is placed at

the point of support rather than anywhere else along the length of wire.


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The placement of additional coils at the point of support in a cantilever does

not alter the maximal elastic load.

A straight wire of a given length and a wire with numerous coils at the point of

support have identical maximal elastic loads, provided they have the same lengths

measured from the force to the point of support.

This should not be surprising since the maximal elastic load is a function of

this length of the configuration rather than the amount of wire incorporated in it. It is

also true for many other configurations: load deflection rate can be lowered without

altering the maximal elastic load if additional wire is properly incorporated. From the

point of view of design, this is important because for the first time, method of

lowering the load deflections rate without subsequently reducing the maximal elastic

load has been discussed.

To achieve this objective with the minimal amount of wire, the optimal

placement of additional wire is at cross-sections where the bending moment is the

greatest. A practical way of deciding where these parts of a wire might be, is to

activate a configuration and see where most of the bending or torsion occurs. These

are the sections where the bending moments or torsion moments are the greatest: the

cross-sections of wire that have the greatest stress.In short it is not the amount of wire

used that is important in achieving a desirably flexible member, but rather it is the

placement of additional wire and its form.

Although additional wire is quite helpful in the design of flexible members of

an orthodontic appliance, it should be avoided in the reactive or rigid members. Loops


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and other types of configurations decrease the rigidity of wire and hence may be

responsible for some loss of control over the anchor units.

STRESS RAISERS

From a theoretical point of view, the force or stress required to permanently

deform a given wire can be calculated; however, in many instances the wire will

deform at values much lower than predicted ones because the presence of certain local

stress raisers increases the stress values in a wire far beyond what might be

predictable by commonly used engineering formulas.

Two common stress raisers are sudden changes in cross-sections and sharp

bends.

A: Any nick in a wire will tend to raise the stress at that cross-section and hence

may be responsible for permanent deformation or fracture at this point. It is

therefore desirable to mark wires by other means than a file, particularly the

wires of smaller cross-sections used in the flexible member of an appliance.

B: A sharp bend in a wire also may result in higher stress than those might be

predicted for a given cross-section of wire. A sudden sharp bend will far more

easily deform than a more rounded or gradual bend. Unfortunately, with a

continuous arch wire, the orthodontist is somewhat limited in space between

brackets and many times is required to make sharp bends because of this

limitation. Flexible member should be designed with gradual bends so that

they will be more free from permanent deformations than comparable ones

with sharp or sudden bends.


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For example, three vertical loops might be compared: a squashed one, a plain

one and one with a helical coil. In terms of permanent deformation, the poorest design

would be loop A, which because of its squashed state has a very sharp bend at its

apex. The plain vertical loop B would be slightly superior, since the bending is more

gradual. Nevertheless a fairly sharp bend occurs at its apex.

The configuration with the most gradual bending is the loop with a helical coil

C. Not only would the helical coil enhance the flexible properties of the spring

because of its additional wire, but the each of gradual bend would further increase its

range of action without permanent deformation.

There are certain sections along a wire where stresses are maximal.

These may be called as critical sections. It has already been seen that in

sections where the bending moments are the largest, areas of high stress exist. These

critical sections are important from the point of view of design, for it is here that

permanent deformation is most likely to occur.

A number of precautions should be observed at critical sections. First stress

raisers should be avoided in these sections at all costs. A nick in a wire, for instance,

might not be so disastrous where the stresses are low, but might will lead to


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deformation or fracture where the stress level is high. Second, the elastic limit of the

wire should be carefully watched at a critical section, lowering the elastic limit at

another place in the wire where the stresses are low, might not be too undesirable but

could be responsible for failure at a critical section.

Therefore in high stress areas it is desirable to use other means of attaching an

auxiliary than soldering or if soldering is to be used as a method of attachment, it

should be done with considerable care.

There are three rules to be kept in mind as far as designs of critical sections.

1) All stress raisers should be eliminated as completely as possible.

2) A large cross-section can be used to strengthen this part of the appliance.

3) The appliance may be so designed that it will elastically rather than

permanently deform under normal loading.

DIRECTION OF LOADING

Not only is the manner of loading important, but the direction in which a

member is loaded can markedly influence its elastic properties. If a straight piece of

wire is bent so that permanent deformation occurs and an attempt is made to increase

the magnitude of the bend, bending in the same direction as had originally been done,

the wire is more resistant to permanent deformation than if an attempt had been made

to bend in the opposite direction. The wire is more resistant to permanent deformation

because certain residual stresses remain in it after the placement of the first bend. If a

bend is made in an orthodontic appliance, the maximal elastic load will not be the

same in all directions. It will be greatest in the direction that is identical to original


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direction of bending or twisting. The phenomenon responsible for this difference is

referred to as BAUSCHINGER EFFECT.

The figure demonstrates a vertical loop with the coil at the apex and a number

of turns in the coil under different directions of loading. The loading in A tends to

wind the coil, increasing the no of turns in the helix and shortening the length. The

type of loading seen in B tends to unwind the helix, reducing the no of coils and

lengthening the spring. The loading in A tends to activate the spring in the same

direction as it was originally wound and hence is the correct method of activation.

ACTIVATION OF HELICAL COIL

A- CORRECT B-INCORRECT

PLACING A REVERSE CURVE OF SPEE


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The same principles can be applied to less complicated configurations such as

in a continuous arch wire. The operator should be sure that the last bend made in an

arch wire is in the same direction as the bending produced during its activation. For

example, if a reverse curve of spee is to be placed in an arch wire, the curve should be

first over bent and than partly removed. Only then will the activation of the arch wire

be in the same direction as the last bend.

FATIGUE OF METALS

Fatigue is the result of repeated stresses at a level, below that which would

normally cause failure. These stresses, usually in the low plastic deformation range,

gradually bring about additional work hardening until the metal finally fails in a

brittle fracture.

Below a certain stress level, a material can be subjected to repeated stresses

without fracture. But fatigue of metal is hastened tremendously by flaws of any kind,

even minute scratch. If there is a defect in the material, such as a scratch or an internal

flaw, the metal remaining around the defect will have to carry an added load and may

lead to failure.

PREVENTION OF FATIGUE FAILURE

Broken wire can add time to treatment. So, it is important that all possible

preventive measures be taken. Care should be taken in wire selection, even though

most suppliers offer wires in which every effort has been made to keep breakage low.

Metals that work hardens rapidly may fatigue more easily. Hard wires are

more brittle than soft wires of the same materials. Hardness level should be selected


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on the basis of individual demands. Experience with specific materials is often the

only criteria in this regard.

During arch designing careful handling should be done. A wire should never

be marked or notched with a file or other sharp instrument. Smooth beaked pliers

should be used to avoid unnecessary damage to the surface, and pliers should be

selected and manipulated so as to avoid marking the wire with the sharp edge of the

beaks.

Smaller diameter wire have a broader working range and may not be so easily

stressed to the proportional limit, as a larger stiffer and seemingly stronger wire. For

this reason change to smaller diameter wire may be the only answer in some cases of

recurrent breakage.

Repeated bending at the same spot should be avoided. All adjustments should

be made away from high stress areas and previous bends at soldered joints should be

avoided, as wire adjacent to solder joints may be subjected to intergranular corrosion

initiated by heat soldering. This can be minimized by careful soldering but additional

protection will be provided by careful cleaning and electro polishing after the

procedure. Good surface finish eliminates many of the small stress raiser that can

initiate the process of failure.


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metal alloys were used routinely for orthodontic purpose because nothing else was

able to tolerate oral conditions.

COMPOSITION

There are two types of Gold wires recognized in American Dental Association

(ADA) specification no 7, year 1984.

Type I: They must contain at least 75% gold and platinum group metals.

Type II: They must contain at least 65% gold and platinum group metals.

In addition to Type I and II Gold wires used in orthodontics before 1950,s two

other types of wires were also used with high content of Gold in at least one of them.

Palladium-Gold-Platinum (P-G-P)

Because of their high fusion temperature and therefore high crystallization

temperature, they are especially useful as wires to be cast against and meet the

composition requirements for an ADA type I wire.

Palladium-Silver-Copper (P-S-C)

These wires are neither Type I nor Type II gold wires, but their mechanical

properties would meet the requirements for an ADA Type I or Type II alloy. The

corrosion resistance of palladium-silver dental alloy, both in cast and wrought forms,

is generally satisfactory.

The basic composition of alloys consists of Gold, platinum, palladium, silver,

copper, nickel and zinc. [Detail in Table]


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WIRE

TYPEGOLD PLATINUM PALLIDUM SILVER COPPER NICKEL ZINC

ADA-I 54-66 7-18 0-8 9-12 10-15 0-2 0-0.6

ADA-II 60-67 0-7 0-10 8-21 10-20 0-6 0-1.7

P-G-P 25-30 40-50 25-30 - 16-17 - -

P-S-C - 0-1 42-44 38-41 16-18 0 -

GENERAL EFFECTS OF THE CONSTITUENTS

1) Gold: Provides Malleability and Ductility.

2) Platinum: It is used to convey greater strength and toughness to assist in

obtaining controllable hardness in the finished wire and contributes

substantially to the resistance of the alloy to tarnish and corrosion by oral

fluids.

3) Palladium: It is the most effective element known for raising, without

widening the melting range of gold alloys. The increased palladium and

platinum content ensures that the wire does not melt or recrystallize during

soldering process. Also these two metals ensure a fine grain structure.

4) Copper: Copper contributes to the ability of the alloy to age harden. When

Copper is present, silver may be added to balance the colour.

5) Nickel: Nickel is sometimes included in small amounts as a strengthener of the

alloy, although it tends to reduce the ductility. The presence of large quantity

of nickel tends to decrease the tarnish resistance and change its response to

age hardening.

6) Zinc: Zinc acts as a scavenger agent to obtain oxide free ingots, from which

the wires are drawn.


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FUSION TEMPERATURE

The minimum fusion temperature of an alloy is usually taken as a temperature

halfway between the liquidus and solidus temperature. Fusion temperature of wrought

wires must be known to ensure that the wires do not melt or lose their wrought

structure during normal soldering procedures.

According to ADA specification no 7, for a type I wire, this temperature is

9550 C (17510 F) or higher, for the type II wire the minimum fusion temperature

should be 8710 C (16000 F).

MECHANICHAL PROPERTIES

Yield Strength Tensile Strength Elongation Fusion Temperature

TYPE MPa 1000psi MPa 1000PSI % % C F

ADATYPE I

582 125 991 117 13 4 995 1750

ADA

TYPE II690 100 862 125 15 2 971 1400

Strength Yield Strength Tensile Strength Elongation Fusion Temperature

P-G-P592-

103480-150

462-

1241125-180 11-15 - 1300-1530

2730-

7750

P-B-C 640-793 100-115965-

1170140-155 16-24 8-15 1050-1080

1710-

1970

A wire of a given composition is generally superior in mechanical properties

to a casting of same composition. The casting contains unavoidable porosity which

has a weakening effect. When the cast ingot is drawn into a wire, the small pores and

surface projections may be collapsed, and welding may occur so that such defects

disappear. Any defects of this type that are not eliminated will weaken the wire.


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PHYSICAL PROPERTIES ARE LISTED IN TABLE

The modulus of elasticity of wrought gold wires is in the range of 97,000 to

117,000 Mpa (14,000,000 to 17,000,000 Psi) which is slightly higher than that for

gold casting alloys. It increases by approximately 5% after a hardening heat treatment.

HEAT TREATEMENT OF GOLD ALLOY

All gold alloy wires that contain copper are heat treatable as the Gold casting

alloys. Type I and II alloys usually do not harden, or they harden to a lesser degree

than do the type III and IV alloys.

The actual mechanism of hardening is probably the result of several different

solid state transformations. Although the precise mechanism may be in doubt, the

criteria for successful hardening are time and temperature.

Alloys that can be hardened, can of course, also be softened. In metallurgic

terminology the softening heat treatment is referred to as solution heat treatment. The

hardening heat treatment is termed as age hardening

SOFTENING HEAT TREATMENT

Gold alloy is placed in an electric furnance for 10 min at a temperature of 7000

C or 12920 F. This is called as annealing. Then it is quenched in water. During this

period all intermediate phases are presumably changed to a disordered solid solution,

and the rapid quenching prevents ordering from occurring during cooling.


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The tensile strength, proportional limit and hardness are reduced by such a

treatment but the ductility is increased.

The softening heat treatment is indicated for structures that are to be ground,

shaped, or otherwise cold worked, either in or out of the mouth. Although 7000 C is

an adequate average softening temperature, each alloy has its optimum temperature

and manufacturer should specify the most favorable temperature and time.

HARDENING HEAT TREATMENT

The age hardening or hardening heat treatment of dental alloys can be

accomplished in several ways. One of the must practical hardening treatments in by “

soaking “ or ageing the alloy at a specific temperature for definite time, usually 15-30

minutes, before it is water quenched. The ageing temperature depends upon the alloy

composition but is generally between 2000 C (4000 F) to 4500 C (8400 F). The proper

time and temperature are specified by the manufacture.

Ideally, before the alloy is given an age-hardening treatment, it should be

subjected to a softening heat treatment to relieve all strain hardening, if it is present,

and to start the hardening treatment with the alloy as a disordered solid solution.

Otherwise, there would not be a proper control on the hardening process, because the

increase in strength, proportional limit, hardness, and the reduction in ductility are

controlled by the amount of solid-state transformations. The transformations in turn,

are controlled by the temperature and time of age-hardening treatment.

COLD WORKING OR WORK HARDENING


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Cold working is also the usual method of hardening gold alloy. Much more

cold working is required for Gold alloys than Steel to harden it. This is to adjust the

drawing and annealing schedule to compensate. Cold working is defined as deforming

a metal at temperature that are low compared with its melting temperatures i.e. any

plastic deformation of metal by hammering, drawing, cold forging, cold rolling or

bending. Gold alloy work hardens much more slowly and to lesser degree than Steel.

To the manufacturer, this low work hardening means that drawing is much easier,

with fewer intermediate anneals required to orthodontist. it means that these metals

are less brittle and will need much more manipulation before they have hardened

excessively.

Some special alloys such as those that are high in platinum, can be harden

materially by temperature manipulation, usually by heating to about 8000 F to 10000 F

and cooling slowly. The slow cooling permits optimum grain growth for the

production of a hard material.


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MICROSTUCTURE

The micro-structural appearance of cold-worked on wrought alloys is fibrous

with extremely elongated crystals. It results from the deformation of the grains during

the drawing operation to form the wire. Such a structure generally exhibits enhanced

mechanical properties as compared with corresponding cast structure. There is a

tendency for wrought alloys to recrystallize during heating operations. The extent of

crystallization is related directly to the duration of heating, the temperature employed,

and the cold work or strain energy imparted to the alloy when the wire was drawn.

Recrystallization is inversely related to the fusion temperature of the wire when

heating temperature and time are constant.

Because there is concomitant decrease in the mechanical properties of alloys

as recrystallization increases, so sufficient platinum and palladium should be present

to increase the fusion temperature of the wrought gold alloy wire. Therefore of all

those wires, the P-G-P wires are the most resistant to recrystallization.

Now a days the use of Gold alloys is markedly reduced because it is too soft to

use as an orthodontic appliance, its high cost, recent advances in the wire materials,

mechanical properties of the same and due to their low yield strength.


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9

Stainless Steel Arch Wires

CARBON STEELS

Stainless steel is the most widely used and accepted material in orthodontics.

It is the major alloy system used in orthodontics. In the mid century stainless steel was

applied to dentistry and orthodontics. Although it was around 1920, that HARRY

BREALY OF SHEFFIELD, F.M.BECKET OF U.S.A. and BENNO STRAUSS

EDWARD MAURS of Germany shared the honor for the development of materials.


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The metallurgy and terminology of these alloys are intimately connected to

those of the simpler binary iron - carbon alloy system and to carbon steel alloys.

Therefore this discussion begins with a brief outline of the metallurgy of the iron-

carbon system.10,26,34,39

Steels are iron based alloys that usually contain less than 1.2% carbon. The

different classes of steel are based on three possible lattice arrangements of iron. Pure

iron at room temperature has a Body Centered Cubic (BCC) structure and is referred

to as FERRITE. This phase is stable at temperatures as high as 9120 C. The spaces

between atoms in the BCC structure are small and oblate; hence, carbon has a very

low solubility in ferrite (maximum of 0.02 Wt %).

At temperatures between 9120 C and 13940 C, the stable form of iron is a Face

Centered Cubic structure (FCC) called AUSTENITE. The interstices in the FCC

lattice are larger than those in the BCC structure. However, the size of the carbon

atom limits the maximum carbon solubility to 2.1 Wt%.

When AUSTENITE is cooled slowly from high temperatures, the excess

carbon that is not soluble in ferrite, forms iron carbide (Fe3C). This hard, brittle phase

adds strength to the relatively soft and ductile ferritic and austenitic forms of iron.

However, this transformation requires diffusion and a definite period of time. If the

AUSTENITE is cooled rapidly (Quenched), it will undergo a spontaneous, diffusion

less transformation to a Body-Centered Tetragonal (BCT) structure called

MARTENSITE. This lattice is highly distorted and strained, resulting in an extremely

hard, strong, brittle alloy.


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The formation of martensite is an important strengthening mechanism for

carbon steels. The cutting edges of carbon steel instruments are ordinarily martensitic,

because the extreme hardness allows for grinding a sharp edge that is retained in use.

Martensite decomposes to form ferrite and carbide. This process can be accelerated by

appropriate heat treatment to reduce the hardness, but this is counter balanced by an

increase in toughness. Such a heat treatment process is called as tempering.

STAINLESS STEELS / CHROMIUM CONTAINING STEELS

When 12 to 30% chromium is added to steel, the alloy is commonly called

stainless steel. Elements other than iron, carbon and chromium may also be present,

resulting in a wide variation in composition and properties of stainless steels.

These steels resist tarnish and corrosion primarily because of the passivating

effect of the chromium. For passivation to occur, a thin, transparent but tough and

impervious oxide layer of Cr 2O3 forms on the surface of the alloy when it is subjected

to an oxidizing atmosphere such as room temperature. This protective oxide layer

prevents further tarnish and corrosion. If the oxide layer is ruptured by mechanical or

chemical means, a temporary loss of protection against corrosion will occur. However,

the passivating oxide layer, eventually forms again in an oxidizing environment.

There are essentially three types of stainless steels, evolving from the possible

lattice arrangement of iron previously described.

TYPE

(SPACE LATTICE)CHROMIUM NICKEL CARBON

Ferratic(BCC) 11.5-27 0 0.20 max

Austantic(FCC) 16.0-26 7-22 0.25 max

Martenstic(BCT) 11.5-17 0-2.5 0.15-1.20


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1. FERRITIC STAINLESS STEELS

These alloys are often designated as American Iron and Steel institute (AISI)

series 400 stainless steels. This series no is shared with the martensitic alloys. The

ferritic alloys provide good corrosion resistance at a low cost, provided that high

strength is not required.

Because temperature change induces no phase change in the solid state, the

alloy is not hardenable by heat treatment. Also, ferritic stainless steel is not readily

work hardenable. This series of alloys finds little application in dentistry.

2. MARTENSITIC STAINLESS STEELS

As noted in above paragraph, martensitic stainless steel alloys share the AISI

400 designation with the ferritic alloys. They can be heat treated in the same manner

as plain carbon steels, with similar results. Because of their strength and hardness,

martensitic stainless steels are used for surgical and cutting instruments.

Corrosion resistance of martensitic stainless steel is less than that of the other

types and is reduced further following a hardening heat treatment. As usual, when the

strength and hardness increases, ductility decreases. It may decrease to as low as 2%

elongation for a high carbon martensitic stainless steel.

3. AUSTENITIC STAINLESS STEELS

The austenitic stainless steel alloys are the most corrosion resistant of the

stainless steels. AISI 302 is the basic type with composition:

Chromium ....... 18%

Nickel ................ 8%

Carbon ........... .15%


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the soldering and welding temperature ranges, normally employee an unavoidable

softening of the wire during normal heating, it is a decided disadvantage.

TYPE

(SPACE LATTICE)

CHROMIUM NICKEL CARBON

Ferratic(BCC) 11.5-27 0 0.20 max

Austantic(FCC) 16.0-26 7-22 0.25 max

Martenstic(BCT) 11.5-17 0-2.5 0.15-1.20

The large modulus of elasticity of stainless steel and its associated high

stiffness necessitate the use of smaller wire for alignment of moderate and severely

displaced teeth. A reduction in wire size results in poorer fit in the bracket and may

cause loss of control during tooth movements. However, high stiffness is

advantageous in resisting deformation caused by extra oral and intra oral tractional

forces.

The yield strength to elastic modulus ratio indicates a lower spring back of

stainless steel than those of newer alloys. The stored energy of activated stainless steel

is substantially less than that of beta titanium and Nitinol wires. This implies that

stainless steel wire produces higher forces that dissipate over shorter periods than

nitinol wires, thus requiring more frequent activation or arch wire changes.

RARK and SHEARER have demonstrated the release of nickel and

chromium from stainless steel appliances.

Low levels of bracket/wire friction have been reported with experiments using

stainless steel wires. This signifies that stainless steel wire offer lower resistance to

tooth movement than other orthodontic alloys.


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HEAT TREATMENT OF AUSTENITIC STEEL

Austenite cannot be hardened like carbon steel by quenching or similar heat

treatment. The only way by which these steels can be hardened is by cold working.

Austenite steel hardens rapidly by cold working with the usual realignment of the

crystalline structure.

Work hardening also brings about some transformation of parts of the

austenite into martensite which adds to the hardening effect.

1. ANNEALING AUSTENITIC STEEL

Stainless steel requires a higher temperature for annealing (18000 F to 20000 F)

than does carbon steel. At this temperature all of the effects of cold working are

eliminated and the metal returns to its softest, most workable state. Orthodontic bands

and ligature wires are usually supplied fully annealed. Cooling from the annealing

temperature must be rapid, usually by quenching. This rapid cooling is not an essential

part of the annealing process, but it is important for corrosion control.

2. STRESS RELIEF OF STAINLESS STEEL

The most important heat treatment process for orthodontic stainless steel is the

relatively low temperature process of stress relieving which is used both in

manufacturing and in orthodontist’s office.

Work hardening steel is hardened by the interlocking of grains and atoms are

locked in situations in which, they are under stress, even when the piece as a whole is

not stressed.


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When a wire with such internal stresses is bend to produce a spring action,

there previously stressed areas can not do their full share.

If the applied force must be resisted by the stressed regions, a part of their

reserve of strength has already been used up by their limit of strength. If the internal

stress is in the same direction as the new load, the two actually augment each other. In

either case, action of the wire is weakened by the internal stress.

Stress relief eliminates such areas of stress within the wire and puts it into the

condition to work most effectively. As internal stresses are relieved, there may also be

some change in the shape of the wire. This is the second reason for stress relieving in

orthodontics. A wire that is bend to form an arch is full of residual stresses which tend

to return it towards its original form. This goes on gradually at ordinary temperature

causing a slow change in arch form (elastic memory). A stress relieving heat

treatment accelerates this change in shape so that the wire will be more stable. When

this treatment is applied to an arch, the form should always be checked and arch

reshaped if necessary after the heat treatment.

Stress relieving changes depend on both time and temperature, and they can be

controlled by the adjustment of either of these factors. In general, low temperature

treatment (4000 F to 7000 F) over a long period of time is most desirable. But, the arch

formed for a patient in the chair cannot be treated for hours or even for too many

minutes. Fortunately, most of the benefits of heat treatment can be produced in few

minutes or less at temperature of 8000 F. This is especially true if the wires have been

previously stress relieved in manufacturing to eliminate the stress in wire making

process.


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The oven is the most reliable method for heat treatment because of relatively

uniform temperature.

INTERGRANULAR CORROSION OF STAINLESS STEEL

Carbon is an undesirable property in austenitic stainless steel, but it is difficult

to remove it completely. The 18-8 stainless steel may lose its resistance to corrosion if

it is heated between 4000 C to 9000C, the exact temperature depending upon carbon

content. Such temperatures are definitely within the range used by the orthodontist in

brazing, soldering and welding.

The reason for decrease in corrosion resistance is the precipitation of

chromium carbide at the grain boundaries at high temperatures. The small rapidly

diffusing carbon atoms migrate to the grain boundaries from all parts of the crystal to

combine with the large, slowly diffusing chromium atoms at the periphery of the

grain, where the energy is highest, and forms chromium carbide (Cr 3C).

The formation of chromium carbide is highest at 6500C. Below this

temperature the diffusion rate is less, whereas, above it, a decomposition of chromium

carbide occurs. When the chromium combines with carbon in this manner, its

passivating qualities are lost, and as a consequence, the corrosion resistance of steel is

reduced.

Because that portion of grain adjacent to grain boundary is generally depleted

to produce chromium carbide, intergranular corrosion occurs, and a partial

disintegration of the metal may result with a general weakening of the structure.

The formation of chromium carbide is called as sensitization.


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