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1 ntroduction 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. 1
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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metallic arch wires are likely to be replaced for most orthodontic applications in the same way as metals have been replaced by composites in aerospace industry.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, 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, dentistry, not only have the 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 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 produce a variety of of stainless steel that would provide low continuous forces over a long long period of time. The NickelTitanium(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 but could still be used in the the 2
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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 NiTi’s perfectly suited this. However, introduction of the TMA 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 individualistic and unique properties 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 4
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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 Nickeltitanium 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
2)
CONVENTIONAL
PSEUDOELASTIC
THERMODYNAMIC
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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T erminology / Definitions MECHANICS
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Is an area of study of physical science, which is concerned with the state of rest or motion of bodies, subjected to forces.
FORCE
Force is defined as an act upon a body that changes or tends to change the state of rest or motion of that particular body.
NEWTON, S FIRST LAW OF MOTION
A particle subjected to a balanced system of concentrated forces will remain at rest, if originally at rest, or will with constant speed in a straight line if originally in motion.
NEWTON, S SECOND LAW OF MOTION
If the particle is subjected to an unbalanced system of forces, the particle will be accelerated in the direction of net force exerted.
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NEWTON, S THIRD LAW OF MOTION
States that, the paired active and reactive forces are equal in magnitude but are directly opposite to one another and are exerted on adjacent particles.
STRESS
Stress is the force per unit area acting on millions of atoms in a given plane of a material OR Displacing forces measured across a given area.
When an external force acts upon a solid body, a reaction force results within the body that is equal in magnitude but opposite in direction to the external force. The external force will be called the load on the body. The internal forces divided by the area over which it acts within the body is the resultant stress. It is measured in terms of pounds/ square inch or psi.
STRAIN
Change in dimension is called as strain. Although strain is dimensionless quantity, units such as m/m or cm/cm are often used to remind one of system of units employed in actual measurements.
Strain may be Elastic Plastic Combination of two
Elastic strain is reversible; it disappears after the strain is removed. Plastic strain is permanent displacement of the atoms inside the material.
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TYPES OF STRESSES AND STRAINS TENSILE STRESS
A tensile stress is caused by a load that tends to stretch or elongate a body. A tensile stress is always accompanied by a tensile strain.
COMPRESSIVE STRESS
If a body is placed under a load that tends to compress or shorten it, the internal resistance to such a load is called compressive stress. A compressive stress is always accompanied by a compressive strain. With both tensile and compressive stress, the forces are applied at right angles to the area over which they act.
SHEAR STRESS
A stress that tends to resist a twisting motion or sliding of one portion of a body over another is a shear on shearing stress. A shear stress is always accompanied by shear strain.
COMPLEX STRESSES
It is extremely difficult to induce a stress of a single type in body. For example, when a wire is stretched, the experimentally observed stress will be predominantly tensile, but the shearing stresses and strain will also be present. Furthermore during
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the deformation, since the volume of wire remains constant, it must decrease slightly in cross-sectional area, a condition that obviously indicates the presence of compressive stresses.
An example of complex stresses as shown in figure is produced by bending a beam in three point loading. As can be seen, compressive, tensile and shear stresses are present in various parts of structure.
ELASTIC LIMIT
If a small tensile stress is induced in a wire, the resulting strain may be such that the wire will return to its original length (i.e. the atoms will move into their regular positions) when the load is removed.
If the load is increased progressively in small increments, and then released after each addition of stress, a stress value finally will be found at which the wire does not return to its original length after it is unloaded. In such a case the wire is said to have been stressed beyond its elastic limit. The elastic limit of a material is the greatest stress to which a material can be subjected, such that it will return to its original dimensions when the forces are released.
PROPORTIONAL LIMIT
If the wire discussed above is loaded in tension in small increments until the wire ruptures without removal of load each time, and if each stress is plotted on a vertical coordinate and corresponding strain is plotted on a horizontal coordinate, a curve is obtained.
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It can be noted that the curve starts as a straight line but gradually curves after a certain stress value is exceeded. If a ruler is laid on a straight line portion of the curve ( from O to P), and if the straight line is extended in a dotted line, the stress at the point P, at which the curve digresses from a straight line, is known as the proportional limit.
HOOKE, S LAW
It states that the stress is directly proportional to the strain in elastic deformation. Since direct proportionality between two quantities is graphically a straight line, the straight line portion of the graph in figure is confirmation of this law.
Since the proportional limit is the greatest stress possible in accordance with this law, it may be defined as the greatest stress that may be produced in a material such that the stress is directly proportional to strain.
YIELD STRENGTH
The yield strength is the stress required to produce the particular offset chosen. The yield strength will always be greater than the elastic limit or proportional limit and will vary with the offset chosen.
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The three terms elastic limit, proportional limit and yield strength are defined differently but their magnitude are so nearly the same that for all practical purpose the terms can often be used interchangeably.
MODULUS OF ELASTICITY
If any stress value equal to or less than the proportional limit is divided by its corresponding strain value, a constant proportionality will result. This constant of proportionality is known as modulus of elasticity or Young’s Modules (E).
Since the modules of elasticity is the ratio of stress to the strain, it follows that, the less the strain for the given stress, the greater will be the value of the modulus. For example, if a wire is difficult to bend, considerable stress must be induced before a notable strain or deformation results. Such a material would posses a comparatively high modulus of elasticity.
The formula for modules of elasticity in tension is derived as follows; Let E = Modules of elasticity F = Applied force on load A = Cross- section of material under stress e= Increase in length l= Original length Stress = F/A =s Strain = e/l = ε Then E = Stress = s Strain = F/A = Fl e/l
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The unit for modules of elasticity is forces per unit area (Mpa or Psi). This property is indirectly related to other mechanical properties.
MAXIMAL FLEXIBILITY
It is defined as the strain that occurs when the material is stressed to its proportional limit. The relation between the maximum flexibility, the proportional limit and modules of elasticity may be expressed as follows: Let
E = Modulus of elasticity P
= Proportional limit
Εm = Maximum Flexibility
From last equation E = P Εm Or Em = P/E
STATIC AND DYNAMIC FORCES
Forces that are applied constantly for an arbitrarily long time are called as static forces / static stresses. The stresses in the teeth during mastication are not of this type. These stresses usually exist for only an instant. They are known as dynamic forces.
Since dynamic forces exits for only very short time, the resulting deformations or strain cannot be measured.
RESILIENCE
It is defined as amount of energy absorbed by a structure when it is stressed, not to exceed its proportional limit. 20
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Properties of Alloys The properties of orthodontic wires are the consequence of two principal origins.
1)
Basic composition will determine the broad range of inherent general properties of each wire type.
2)
The particular nature of the drawing process including the heat treatment by the manufacturers and clinician will have further significant effects on the specific properties.
BASIC PROPERTIES OF ELASTIC MATERIALS
3,18,19,22,24
The elastic behavior of any material is defined in terms of its stress – strain response to an external load. Both stress and strain refer to the internal state of the material being studied: stress is the interval distribution of the load, defined as force
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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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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 23
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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 29
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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 treatment is applied. applied. Each grain is elongated elongated in the same proportion proportion as the Ingot. Ingot. The squeezing, squeezing, massaging massaging action action of rolling the Ingot 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 b y 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 qualit y 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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RECTANGULAR WIRES
Rectangular wire can be made by drawing the materials through a rectangular die or by rolling round wires to a rectangular r ectangular shape. There appears to be no significant difference in the wires formed by the two processes but is difficult to evaluate. Round wires made by drawing, vary as much in physical properties as most of rectangular wires. Therefore it would be unrealistic to attribute specific differences to rolling and drawing process.
Drawing can however produce a sharper corner on a rectangular wire and this can be an advantage in the application of torque.
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7 deal Orthodontic Alloy I deal
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 asticity (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 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 (d 3). In the design of the active member it is good policy to use as small as cross – cross – section 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 LOAD DEFILATION RATE
MAXIMUM INCREASE
MAXIMUM DEFLECTION
Activation of wire without changing length
decreased
No change
Increase
Activation in direction of original bending
-
Increase
Increase
If rate is maintained as constant
Increase as 1/h
Increase as 1/h
MECHANICAL PROPERTIES OF WIRE
MODULUS OF ELASTICITY
PROPORTIONAL LIMIT
SP/E
Cross section(round)
L
d
1/d
Cross section(rectangle)
h
h
1/h
1/L
1/L
L
DESIGN FACTOR
Alteration of cross section to rectangular form
Length/cantilever
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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 crosssection. 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
150.06
0.016
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 Cross section Shape
CS FIRST ORDER 130.52
SECOND ORDER 132.63
1129.79
297.57
0.457X0.035 0.535X0.035
1805.10 2173.95
966.87 1535.35
0.546X0.711
3129.83
1845.37
M
mm
RECTANGULAR RECTANGULAR
0.010 X 0.020 0.016X0.022
0.254X0.508 0.406X0.550
RECTANGULAR RECTANGULAR
0.018X0.025 0.021X0.025
RECTANGULAR
0.0215X0.028
Cross section
Shape
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):
Where
S
= Ws x As
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 42
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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 49
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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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8
Gold Wires Pure Gold is the noblest of all dental metals, rarely tarnishing and corroding in the oral cavity. It is inactive chemically, and it is not affected by air, heat, moisture and most solvents. It is the most ductile of all metals, as demonstrated by its ability for a 1oz cylinder to be drawn into a wire 100 km long in length. It is the most malleable of all metals, as shown by its ability to be hammered to a thickness of 0.00013 mm, about one third of thinnest gold foil used in dentistry. 34,39
Pure gold is extremely soft, but after cold working, its hardness
( 52 to 75
Vickers hardness no [VHH]) is equivalent to and may exceed that of conventional Type I gold alloy (50 VHN) in its softened state, and after work hardening, its hardness approaches that of Type II Gold alloy (90 VHN).
Although its ductility decreases after cold working, wire is the principal form in which wrought gold dental alloy is used. Before 1950,s Gold and other precious
53
D#$5 E'-18
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]
54
D#$5 E'-18
WIRE TYPE
GOLD
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.
55
D#$5 E'-18
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 871 0 C (16000 F).
MECHANICHAL PROPERTIES Yield Strength
Tensile Strength
Elongation
Fusion Temperature
TYPE
MPa
1000psi
MPa
1000PSI
%
%
C
F
ADA TYPE I
582
125
991
117
13
4
995
1750
ADA TYPE II
690
100
862
125
15
2
971
1400
Strength
Yield Strength
Tensile Strength
Elongation
Fusion Temperature
P-G-P
5921034
80-150
4621241
125-180
11-15
-
1300-1530
27307750
P-B-C
640-793
100-115
9651170
140-155
16-24
8-15
1050-1080
17101970
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 700 0 C or 1292 0 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 ti me.
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 200 0 C (4000 F) to 450 0 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 800 0 F to 1000 0 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 thei r 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. 61
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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 ironcarbon 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 912 0 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 1394 0 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 (Fe 3C). 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% 64
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Type 304 stainless steel has a similar composition, but the chief difference is its reduced carbon content (0.08%).
Both 302 and 304 stainless steel may be designated as 18-8 stainless steel. They are the types most commonly used by the orthodontist in the form of bands and wires.
Generally austenitic stainless steel is preferable to ferritic stainless steel because of the following characteristics: Greater ductility and ability to undergo cold work without fracturing. Substantial strengthening during cold working. Greater ease of welding. Ability to fairly and readily overcome sensitization. Less critical grain growth. Comparative ease in forming.
MECHANICAL PROPERTIES
The property of readily strain hardened is a characteristic of austenitic stainless steel. Part of this increase in hardness is ordinary strain hardening. But a considerable amount is the result of phase change from a face centered to a body centered lattice. This phase change can be readily demonstrated, since the body centered lattice are ferromagnetic at room temperature, austenitic is non magnetic. It is unfortunate that after strain hardening, a stainless steel wire can become fully annealed in few seconds at a temperature of 7000 C to 800 0 C. After such an annealing, it has lost much of the range of elasticity or working range, so necessary to a satisfactory orthodontic appliance. Because the annealing temperature involved in
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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
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
(SPACE LATTICE)
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 (1800 0 F to 2000 0 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 700 0 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 800 0 F. This is especially true if the wires have been previously stress relieved in manufacturing to eliminate the stress in wire making process. 68
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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 400 0 C to 900 0C, 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 650 0C. 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. 69
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PREVENTION OF INTERGRANULAR CORROSION
There are several methods by which this condition can be minimized. Two most commonly used methods are:
1.
Keeping out of the sensitizing temperature range.
Speed in handling the metals in the sensitizing temperature range, as during soldering can be very effective means of minimizing sensitization.
Stainless steel should always be quenched immediately after soldering to bring it down to a safe temperature as soon as possible. This is also the reason for quenching after annealing. At annealing temperature the chromium carbide is broken up. If a metal is cooled rapidly from annealing to room temperature there is no opportunity for chromium carbide to form.
Both low temperature and high temperature solders can be used to control intergranular corrosion. If they are used properly with low temperature solder (silver solder that melts below 1100 0 F) the objective is to heat it to soldering temperature, solder, and then quench as quickly as possible. This is the most commonly used procedure in soldering stainless steel. High temperature solder (Gold solder that melts above 1200 0F ) also can be used, but only if the entire piece of steel can be heated to this high temperature. The metal is then above the sensitizing range, while it is being soldered, and thus it is perfectly safe. Of course, it must be quenched immediately after soldering. If only part of the steel is heated to this high soldering temperature, there will be a zone
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outside the soldering area which is in the sensitizing temperature range. Therefore this method is useful only for small pieces.
2.
Controlling the carbon content (stabilization)
The second method for control of intergranular corrosion is introduction of some elements which tie up with chromium or by keeping the carbon content exceptionally low (below .08%). Titanium is often used for this purpose.
If titanium is introduced in an amount approximately six times the carbon content, the precipitation of chromium carbide can be inhibited for a short period at the temperatures ordinarily encountered in soldering procedures. Stainless steel that have been treated in this manner are said to be stabilized. Stabilized steel is less susceptible to intergranuler corrosion but it is still not 100% safe. Proper handling by orthodontist can modify the advantage completely.
BRAIDED, TWISTED OR MULTISTRANDED WIRES
Very small diameter stainless steel wires can be braided or twisted together by the manufacturer to form larger wires for clinical orthodontics. In terms of performance, the wire is delivering higher forces per unit of activation over a greater distance and strength is also increased. The result is an inherently high elastic modulus material with low stiffness because of its co-axial spring like nature.
The separate strands may be as small as 0.178 mm, but the final intertwined wires may be either round or rectangular in shape, and their cross-sectional dimension is in between .406 mm and .635 mm.
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Because of their low “apparent” elastic modulus in bending, these types of wires apply low forces for a given deflection when compared to solid stainless steel wires.
Kusy and Dilley investigated the strength, stiffness and springback properties of multistranded wires in a bending mode of stress. They noted that the stiffness of a triple stranded 0.0175″ (3 x 0.008″) stainless steel arch wire was similar to that of 0.010″ single stranded stainless steel wire. The 0.0175″multi stranded wire was 25% stronger than 0.010″ stainless steel wire.
The 0.0175″ multi stranded wire and 0.016″ nitinol showed similar stiff nesses. However Nitinol tolerated 50% greater activation than the multi stranded wires. The triple stranded wire was also half as stiff as a 0.016″ beta titanium wire. 26
Ingram, Gipe and Smith noted that titanium alloy wires and multi stranded stainless steel wires have low stiffness when compared with solid stainless steel wires.
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The investigators also found that most multi stranded wires had a spring back similar to that of nitinol, but a larger spring back when compared with solid stainless steel or beta-titanium wires and they have spring back properties that are relatively independent of wire size.
ADVANTAGES OF STAINLESS STEEL
Lowest cost of the wire alloys. Proven biocompatibility from extensive clinical use Excellent formability for fabrication into orthodontic appliances. Can be soldered and welded, although welded joints may require solder reinforcement.
DISADVANTAGES OF STAINLESS STEEL
High force delivery Relatively low spring back in bending compared to beta-titanium and Nickel titanium alloys. Can be susceptible to intergranular corrosion after heating to temperatures required for joining.
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111
Chrome-Cobalt -Nickel Alloy Archwire Cobalt – Chromium – Nickel Orthodontic wires are very similar in appearance, mechanical properties and joining characteristics to stainless steel wires, but have a much different composition and considerably greater heat treatment response.34,39 These alloys were originally developed for use as watch springs (ELGILOY), but their properties are also excellent for orthodontic purpose. These wires are available in four tempers: soft, ductile, semi resilient and resilient. The differences in mechanical properties arise from proprietary variations in the wire manufacturing process.
The wires are furnished to orthodontist in different gauges and cross-sectional shapes with differing physical properties. Their resistance to tarnish and corrosion in the mouth is excellent. They can be subjected to same welding and soldering procedures as described for stainless steel orthodontic wires.
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COMPOSITION:
COBALT
40%
CHROMIUM
20%
NICKEL
15%
MOLYBDENUM
7%
MANGANESE
2%
CARBON
.15%
BERYLLIUM
.04%
IRON
15.8%
The soft temper wires are popular with clinicians because they are easily deformed and shaped into appliances, then heat treated to provide substantially increased values of yield strength and resilience. The other tempers are less popular than the soft temper because wires made from them have lower formability and are somewhat higher in cost than stainless steel
HEAT TREATMENT
Cobalt-Chromium-Nickel alloy may be softened by heat soaking at 1100 0C to 1200 0C, followed by a rapid quench. The age hardening temperature range i s 260 0C to 6500C. For the alloy Elgiloy the alloy should be held at 482 0C for 5 hours.
Ordinarily, the wires are heat treated before supplied to the user and may be ordered in several degrees of hardness. In addition, the orthodontist can heat treat the wires by placing them in an oven or by passing an electric current through them with certain types of spot welders. A typical cycle would be 482 0C for 7 to 12 minutes. This heat treatment would increase the yield strength and decrease the ductility.
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Alloys
Modulus of Elasticity 3 (10 MPa [GPa]
0.2% Offset yield strength (MPa)
Ultimate Tensile Strength
Number of 90 degree cold bends without fracture
S.S
179
1579
2117
5
Co-Cr
184
1413
1682
8
NiTi
41.4
427
1489
2
B-TITANUM
71.7
931
1276
4
Wires made from this alloy should not be annealed. The resulting softening effect cannot be reversed by subsequent heat treatment. Moreover, if only a portion of a wire is annealed, severe embrittlement of adjacent sections may occur.
PHYSICAL PROPERTIES
Tarnish and corrosion resistance are excellent. Hardness, yield strength and tensile strength are approximately the same as those of 18-8 stainless steel. Typical mechanical properties of orthodontic wires are shown in table. Ductility in the softened condition is greater than that of 18-8 stainless steel alloys and less than the alloys in the hardened condition.
RECOVERY HEAT TREATMENT
An increase in the measured elastic properties of a wire can be affected by heating it to comparatively low temperatures (370 0C to 480 0C) after it has been cold worked. The stress relief heat treatment removes residual stresses during recovery without pronounced alteration in mechanical properties. Such a treatment also stabilizes the shape of the appliance.
Cobalt-Chromium-Nickel wires are more responsive than the 18-8 stainless steel wires to the low temperature heat treatment. A reduction in ductility
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accompanies the increase in yield strength. A phase change as well as stress relief is probably responsible. Caution must be used to avoid excessive embrittlement.
Although the optimum temperature range for the stress relief heat treatment is most often reported at 370 0C to 480 0C, there appears to be no reason to exceed the low temperature limit of 370 0C when the wire is in nonstabilized grade of austenitic stainless steel. Eleven minutes at approximately 3700 C results in a maximum proportional limit for a severely cold worked appliance. This temperature is also below the lower limit (4250 C) of the sensitization temperature range.
The softest Elgiloy (Blue) cannot be heat treated to become as brittle or hard as the high spring temper, or hardest Elgiloy (Red) wire. To use Elgiloy properly the user must be thoroughly familiar with it. A stress relief heat treatment not only improves the working elastic properties of a wire appliance but also can reduce failure caused by corrosion, which may occur in areas of high localized s tress.
MECHANICAL PROPERTIES
With the exception of red temper Elgiloy, non heat treated co-cr wire have a smaller spring back than stainless steel wires of comparable size, but this property can be improved by adequate heat treatment. Optimum levels of heat treatment can be confirmed by a dark straw colored wire or by use of temperature indication paste.
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The advantage of co-cr wire over stainless steel wires includes greater resistance to fatigue and distortion and longer function as a resilient spring. In most other respects, the mechanical properties of co-cr wires are very similar to those of stainless steel wires. Therefore stainless steel wires may be used instead of co-cr wires of same size in clinical situations in which heat hardening capability and added torsional strength of co-cr wires are not required.The high modulus of elasticity of cocr wires suggest that these wire deliver twice the force of Beta-titanium wires and four times the force of Nitinol wires for equal amount of activation.
Co-cr wires have good formability and can be bent into many configurations relatively easily. Caution should be exercised when soldering attachments to these wires, since high temperature can cause annealing with resultant loss in yield and tensile strength. Low fusing solder is recommended for this purpose.
ADVANTAGES
Relatively low cost, although greater than stainless steel Proven biocompatibility from extensive clinical use Outstanding formability in as-received condition. Can be soldered and welded. Excellent corrosion resistance in mouth.
DISADVANTAGES
High elastic force delivery Lower spring back than stainless steel.
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RECENT ADVANCES IN COBALT-CHROMIUM 1.
G and H WIRE COMPANY
Combining ductility and strength, colboloy, nickel-cobalt wires can be heat treated in bend areas and easily soldered without annealing. They are highly flexible and resistant to set, fatigue and corrosion. The wire also offers reduced bracket friction and greater spring efficiency than typical stainless steel wires.
A true arch form is available in sizes from 0.014″ to 0.018″ round and 0.016″ x 0.022″ to 0.019″ x 0.025″ rectangular.
2.
MASAL ORTHODONTICS INTERNATIONAL
Heat treatable blue Masiloy chrome-cobalt arches can accept sharp bends without breakage. Heat treatment increases the resiliency by 20%. The wires are available in natural arch sizes of 0.016″ x 0.016″, 0.016″ x 0.022″ and 0.017″ x 0.025″.
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11
Australian Archwires
In collaboration with an Australian metallurgist, Begg sought to develop a wire material which met his paradoxical requirements. After several years of experimentation they produced a wire which is thus enough to distribute forces at an optimal level for tooth movement over a considerable distance for a long period of time and with a minimal loss of force and intensity while doing so. It is also thick enough to resist weakening and distortion due to the wear and tear exerted on appliances within the mouth. 4,38 80
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In addition to variables like wire drawing and heat treatment, slight fluctuation in the speed at which the wire is drawn through the dies affects its physical properties. Additional variations can be caused by fluctuations in the rate at which the wire passes the heat source, and this may combine to aggravate the variations introduced previously.
This quality control problem may result in a wire that is too soft or too brittle. Another problem is that the wire may seem satisfactory, yet contains flaws which cause breakage during arch fabrication.
One of the outstanding property of Australian wire is its resilience or ability to spring back after having been deflected. This property can be checked by bending the wire with the fingers while holding it with the pliers.
Australian wires are available in the following forms. 1.
Regular grade
White
2.
Regular plus
Green
3.
Special
Black
4.
Special plus
Orange
5.
Extra special plus
Blue
6.
Supreme
Blue
REGULAR GRADE
Lowest grade and easiest to bend. Used for forming auxiliaries and can be used for forming arch wires when distortion and bite opening is not a problem.
Available in sizes of 0.012″, 0.014″, 0.016″, 0.018″ and 0.020″
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REGULAR PLUS GRADE
Relatively easy to form, yet more resilient than regular grade. Used for auxiliaries and arch wires when more pressure and resistance to deformation is desired. Available in sizes 0.014″, 0.016″, 0.018″ and 0.020″.
SPECIAL GRADE
Highly resilient, yet can be formed into intricate shapes with little danger of breakage. The 0.016″ is often used for starting arches. Available in sizes 0.014″, 0.016″, 0.018″ and 0.020″.
SPECIAL PLUS GRADE
Special plus wire is routinely used by experienced operators. Hardness and resiliency of 0.016″ wire is excellent for supporting anchorage and reducing deep over bites. Available in sizes 0.014″, 0.016″, 0.018″, 0.020″ and 0.022″.
EXTRA SPECIAL PLUS
This grade is unequated in resilience. It is more difficult to bend and more subject to fracture. However many orthodontists feel that the ability of this wire to move teeth, open deep overbites and resist deformation far outweighs the inconvenience caused by an occasional breakage while bending. This wire can be easily broken if not bent properly; there is no margin for bending errors. Each 25 foot spool is pretested, an intermaxillary hook is bent in the wire and left as evidence that the wire is of proper quality.
Available in size of 0.016″only.
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SUPREME GRADE / PREMIUM PLUS
Primarily used only in treatment of rotations, alignment and leveling. It is intended for use in either short sections or full arches where sharp bends are not required.
Available in 0.010″, 0.012″ and 0.016″.
Due to extreme hardness of Australian wire, special attention must be given to bend it successfully.
1.
Pre-warm the wire by sliding between the thumb and forefinger. Do not attempt to straighten the wire by stripping between the plier b eaks.
2.
Hold pliers very lightly when bending the wire. Do not squeeze or pull the wire. Pliers must have smooth beaks, carbide tips are not recommended.
3.
Bend the wire very slowly pressing with the thumb or forefinger. Do not rotate the pliers while beading loops and circles should be formed against the square beak and beaks should be apart slightly.
4.
Never pinch the wire with the pliers before or during bending.
5.
Do not scratch the wire to locate bends.
Australian wires become hard from bending (work hardening). Hence, there is no need for heat treatment and no margin for back bending to correct mistakes.
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USE OF NEWER AUSTRALIAN WIRES
The supreme grade of sizes 0.008″- 0.011″ are used for -
Relieving
crowding.-For
making
different
auxiliaries
like
MAA
(Mollenhauer,s aligning auxiliary), Spec auxiliary, Udder arch etc. -
For making mini uprighting springs.
-
0.011" wire can be used for aligning second molar towards the end of stage III.
MANUFACTURING OF A.J. WILCOCK ARCH WIRES SPINNER STRAIGHTENING
It is a mechanical process of straightening resistant materials, usually in the cold drawn condition. The wire is pulled through rotating bronze rollers which torsionally twist the wire into straight condition. The disadvantage of this process i s
-
Resultant deformation
-
Decrease yield stress value.
PULSE STRAIGHTENING
In pulsed straightening, the wire is pulsed in a special machine which permits high tensile wires to be straightened and into lower diameters than possible earlier with spinner straightening. The material yield strength is not altered and surface has a smoother finish and therefore causes low friction.
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12
N ickel Titanium Arch Wires
A significant advancement in orthodontic materials was made in the late 1930,s and 1940,s when stainless steel wire became widely available. Since that time there has been continuous evolutionary improvement in the strength and resiliency of wires used for orthodontic treatment. The development of Nitinol wire was another improvement which emerged from the orthodontic search for lighter force and greater working range.1,6,7,10,11,13,14,15,16,17,20,33,34,35,37,39
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Titanium, a metal discovered by M.H.KLAPROTH in 1795, has changed very rapidly from a rare metal to an important structural metal because of its high weight, high strength and corrosion resistance. It has an atomic number 22 and atomic weight 47.9 and occupies ninth place in abundance of metals in earth’s crust. 98% of all rocks examined contained titanium besides sand, clay and other soils. Many minerals contain titanium, the main ones being Ilmenite and Rutile. Ilmenite is a non-titanium oxide ore or Iron titanate which contains 32% titanium. Rutile is titanium oxide and is richer in titanium content.
Nitinol was invented in early 1960,s by WLLIAM F. BUEHLER, a research metallurgist at the Naval Ordinance Laboratory in Silver Spring, Maryland. (Now called as Naval Surface Weapons Centre). He did extensive research and published his findings on the properties and uses of this new alloy.
The name Nitinol is an acronym derived from the elements which comprises the alloy, Ni from nickel, Ti from titanium and Nol from Naval ordinance l aboratory.
CONVENTIONAL NITINOL
Niti was introduced to orthodontics by Dr. GEORGE ANDREASEN and his associates. Largely through his efforts and those of the Unitek Company, the first Nitinol alloy was marketed to orthodontists as Nitinol. Ironically, this first 50:50 composition of Nickel and titanium was a shape memory alloy in composition only.
Indeed this alloy was passive, as the shape memory effect had been suppressed by cold working the wire during drawing. What was so attractive about this martensitic stabilized alloy was its low force per unit of deactivation, i.e., its low
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stiffness. Compared with the competition of the day, this wire was quite springy, delivering only 1/5 th to 1/6 th the force per unit of deactivation and better meeting the criteria of light, continuous force. When this stiffness was combined with its outstanding range and high spring back, one might presume that this wire was the ideal. It did not take long, however, before its lack of formidability was recognized as limitation, especially when wires broke.
The lack of formability largely remains today, but the initial brittleness that plagued the early nitinol product has long since been rectified.
PSEUDOELASTIC NITINOL
In addition to conventional Martensitic stabilized alloy, two other generic nitinol type alloys are available today that are active, i.e. they undergo some form of shape memory effect (SME) and are super elastic. Two generic alloys are 1.
Austenitic active alloy
2.
Martensitic active alloy
In the austenitic active alloy, both the martensitic and austenitic phases play an important role during its mechanical deformation.
Martensite represents the low
stiffness phase and austenite represents the high stiffness phase. Thus on loading, the austenitic active alloy produces some three times the force per activation of the conventional martensitic stabilized nitinol alloy. Fortunately this effect is short lived. At first glance one would suspect that the mechanical properties are dismal, but presenvence prevails, wherein the stiffness is comparable to that of martensitic
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nitinol. In fact a stress induced phase transformation has occurred in which the austenitic phase has transformed to the martensitic phase. Upon deactivation the reverse occurs and martensitic phase is gradually transformed to the austenitic phase. Because the spring back is nearly total, this series of clinical events is elastic despite the fact that the appearance is quite non linear. Here the martensite reversibly transforms to austenite and thereby changes shape to maintain force, represent the key attribute to this nonlinear but nonetheless elastic alloy and is called pseudo elasticity. Today several alloys are being marketed that utilize pseudo elasticity. Most common of these is 270C super elastic copper niti described later.
THERMOELASTIC NITINOL
The third Nitinol type alloy in the market today is a martensitic active alloy that ultimately exhibits a thermally induced Shape Memory Effect (SME). This is the long awaited Nitinol alloy that Dr. ANDREASEN hoped to someday employ in orthodontics. For many years the alloy compositions simply could not be controlled precisely enough to make a uniform wire product. Transition temperature from Martensitic to Austenitic had to occur in the region of ambient oral temperature, and yet it was known that for every 150 parts per million variations in composition, a 1 oC change in transition temperature occurred. After a 20 years delay, MIURA showed that surgical cases could be treated by preparing a series of arches in which the desired shape was set by heat. Upon distortion and insertion into patient’s mouth, the appliance would be activated by the warmth of oral cavity and return to its predetermined shape. By capitalizing on thermo elasticity, a series of final arch forms could be generated and thereby the practitioner could maintain control. Today the thermo elastic effect is demonstrated in GAC international’s alloy, Sentalloy light. 88
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COMPOSITION AND PHYSICAL PROPERTIES
The composition of Nickel-Titanium used in dentistry is as follows: NICKEL
-
54%
TITANIUM
-
44%
COBALT
-
2%
This composition results in a 1 to 1 atomic ratio of the major components. As with other systems, this alloy can exist in various crystallographic forms.
At high temperatures a body centered cubic lattice (BCC), referred to as austenitic phase is stable, whereas appropriate cooling can induce transformation to a close packed hexagonal martensitic lattice. This characteristic transformation of austenitic to martensitic phase results in two unique features of potential clinical relevance i.e. shape memory and super elasticity or pseudo elasticity.
SHAPE MEMORY WIRE
Most orthodontists are aware of Nitinol because of unique property of the alloy called “shape memory”. Nitinol has the characteristics of being able to return to a previously manufactured shape when it is heated through a transition temperature range. To use this property, the wire must first be set into the desired shape and held while undergoing a high-temperature heat treatment. After the wire has cooled to room temperature, it may be deformed within certain strain limits.When heated to its unique transition temperature range, it will remember its shape and return to the original configuration. Though the orthodontic wires available today do not fully utilize this characteristic, research into orthodontic application for the “memory” aspects of Nitinol wire are continuing at the University Of Iowa and at unitek. 89
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The “memory” effect is achieved by first establishing a shape at temperature near 4820C (9000 F). The wire is then cooled and formed into a second shape. Subsequent heating through a lower transition temperature causes the wire to return to its original shape. The cobalt content is used to control the lower transition temperature which can be near mouth temperature 37 0C (98.40 F).
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SUPER ELASTICITY
Inducing the austenitic to martensitic transition by stress can produce super elasticity, a phenomenon which is employed with some Nickel – orthodontic wires. If alloy is stressed, it initially results in the strandard proportional stress-strain behavior. However at a stress, sufficient to induce phase transformation, there is sufficient increase in strain, referred to as super elasticity or pseudo elasticity.
This additional strain is due to the volume change that results from change in crystal structure.
Nickel titanium alloy, therefore can be produced with either the austenitic or mastensitic structure having varying degree of cold work and variations in transition temperature. In general Nickel-titanium wire has relatively low modulus values and larger working range. They are difficult to join and have to be joined by mechanical crimps, since the alloy can neither be soldered nor welded.
OTHER PROPERTIES
Clinch back distal to molar buccal tube can be obtained by resistance or flame annealing the end of the wire. This makes the wire dead soft and can be bent into preferred configuration. A dark blue color indicates the desired annealing temperature. Care should be taken not to overheat the wire because this makes it brittle.
Findings on resistance to corrosion of nitinol wires have been inconsistent. Although some investigators reports that nitinol is resistant to corrosion as stainless steel, various authors have found nitinol to be more susceptible to corrosion than other
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orthodontic alloys. SCHWANINGER, SARKAR and FOSTER have noted that corrosion does not affect flexibility properties of nitinol wires. Some reports indicates an increase in permanent deformation and decrease in elasticity caused by corrosion or the cumulative effects of cold working.
Nitinol has been reported to be more susceptible to electrolytic dissolution than stainless steel. The examination of unused wires revealed large variation in surface texture of nitinol wire when compared to stainless steel. With small metallic prominences Nitinol wire frequently exhibits an undulating bubbling or mottled cake appearance. The electrolytically corroded Nitinol wires have either obvious pits which occurs along the sharp edges of rectangular wires where the electric field would be greatest or have very irregular surfaces with loosely bound corrosion products.
Used unclean wires are frequently covered with organic layer which possesses significant elements such as Na, P, S, Cl, K and Ca. These layers are not usually present on cleaned surfaces. Because of the variation of unused Nitinol surface and possible organic contamination, it is difficult to assess the degree of corrosion existing on surfaces by visual inspection.
The most important benefit from Nitinol wires are realized when a rectangular wire is inserted early in treatment. Simultaneous rotation, leveling, tipping and torquing can be achieved with a resilient rectangular wire such as nitinol. Clinicians have been successful in beginning treatment of certain carefully selected cases with full size rectangular wires that nearly fills the bracket slot. In few instances the active case has been treated with just one arch wire.
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It is ideally suited for use with most pretorque and preangulated appliances because tipping and up righting of the teeth can be initiated in the early stages of treatment. When the case is nearing completion with a nitinol arch wire, there is very little to be done in the way of placing compensating bends to upright roots, once the spaces have been closed. In the treatment of extraction cases with pretorqued and preangulated twin brackets and Nitinol, conventional auxiliary method of closing spaces may be used along with headgear when needed. The use of nitinol with pretorqued and preangulated brackets, require careful monitoring of tooth movements because of wires high elasticity and more continuous force. Therefore time intervals between appointments cannot be extended.
MECHANICAL PROPERTIES
The modulus of elasticity of nitinol is 41.4 x 10 3 Mpa, (6 x 10 6 psi), the yield strength is 427 Mpa (62,000 psi), and the ultimate tensile strength is 1489 Mpa (216,000 psi). These properties results in very low orthodontic forces when compared with similarly constructed and activated stainless steel. The low stiffness in combination with moderately high strength accounts for wire’s large elastic deflection or working range. The alloy has limited formability.
Elasticity (10 MPa [GPa]
0.2% Offset yield strength (MPa)
Ultimate Tensile Strength
Number of 90 degree cold bends without fracture
S.S
179
1579
2117
5
Co-Cr
184
1413
1682
8
NiTi
41.4
427
1489
2
B-TITANUM
71.7
931
1276
4
Modulus of Alloys
3
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CLINICAL APPLICATION OF NITI
High spring back, flexibility, low constant forces, shape memory and elasticity are important and advantageous properties for clinical application of Niti.
GARHER, ALLAI, MOORE and KAPILA and associates have noted that bracket/wire frictional forces with nitinol wires are higher than those with stainless steel wire and lower than those with beta-titanium.
Niti can be successfully used in the treatment of Cross bite corrections Up righting impacted canines Opening the bites.
Nitinol wires can be used in Class I, Class II or class III malocclusions, in both extraction on non extraction cases. In selecting cases that benefit most from the use of nitinol wires, the primary criteria is the amount of malalignment of the teeth from the ideal arch form. The more the wire has to be deflected from the ideal arch form when ligated into the bracket, the greater benefit nitinol wire has over stainless steel.
USE OF THERMOELASTIC NITINOL
Nitinol has unique property which is of practical use to the orthodontist. That property is its extreme elasticity when it is drawn into high-strength wire. This wire is much more difficult to deform during handling and seating in brackets slot than stainless steel. It’s the nitinol extreme elasticity that offers the clinician an advancement in the application of orthodontic materials.
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The wire can be used for a longer period of time without changing and it can shorter the treatment time needed in leveling the dentition. Nitinol has another remarkable property of returning to a previously manufactured shape, when it is heated through a transition temperature range. If we were to take advantage of this property, the wire must first be set into the desired shape while undergoing a high temperature heat treatment. After the wire has cooled to room temperature, it may be deformed within certain strain limits. When heated to its unique TTR it will remember its shape and return to the original configuration.
Nitinol wire after being deformed will spring back to its original shape by either of two methods. First, it will experience a nearly complete spring back because of its modulus of elasticity without heat, restoring the desired wire to its original shape.
Second it will experience a complete spring back from the deformed shape by being placed in the TTR between 31 to 45 oC. The average temperature of the mouth is in this range and triggers the wire to assume the original shape bent into it.
The TTR of Nitinol can be adjusted by varying the Nickel and Cobalt content. For orthodontic purpose, this thermal nitinol wire is alloyed so that the TTR corresponds to the approximate temperature in the mouth and therefore allows part of the wire “memory” property to be used for moving teeth.
CLINICAL RECYCLING OF NICKEL TITANIUM
Now a days two types of Nickel titanium alloy wires are commercially available. First of them is Nitinol (Unitek Corporation) and the recent one marked as Niti Cormaco (Alif) and Sentalloy (GAC International) among others. The two
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demonstrates several differences in properties. Whereas the original Nitinol wires are primarily in the martensitic phase at room temperature, the newer Niti wires have an austenitic grain structure and 1.6 times greater spring back. When compared with Nitinol, the Niti wires are also 36% stiff at 80% of activations and are not time dependent with regard to stress relaxation.
The desirable mechanical properties of Nickel Titanium alloy wires and their relatively high cost has prompted many clinicians to recycle these wires.
Nickel Titanium wires, especially those of newer pseudo elastic type, undergoes phase changes as a result of heat treatment that substantially alter their properties. KAPILA, HAUGEN and WATANABE noted that temperatures greater than 600C increases the susceptibility of these newer austenitic Nickel Titanium wires to plastic deformation and decreases their spring back. Since various forms of heat treatment are often used for sterilization, further studies to determine the effects of recycling in conjuncture with heat sterilizations are indicated.
STERELIZATION OF ORTHODONTIC ARCH WIRES
The relatively high cost of nickel titanium wires and it ability to return to its original forms had raised concern about the treatment of the wire between patients for prevention of cross- infection.
BUCKTHAL and KUSY had studied the effects of cold disinfectants on the mechanical properties and surface topography for 0.17″ x .025″ Nitinol wires. Three disinfections approved by the ADA were used at maximum antimicrobial concentration. 2% acidic glutaraldehyde, chlorine di oxide and iodophor were used. Bending and
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tensile tests were conducted to determine weather the stiffness, strength or range of the wires changes after disinfectant treatment. Further surface topography was also studied with laser spectroscopy to see surface changes from tarnish and corrosion.
The result showed no significant changes in fundamental stiffness or inherent strength of the wires after multiple disinfectant cycles. Wires showed no additional surface pitting or corrosion.
STAGGERS and MARGESON studied the effects of various types of sterilization on tensile strength of orthodontic wires. The sterilization methods investigated were Dry heat using the Dentronix DDS 5000 dry heat sterilizer (375 0 F for 20 min), Autoclaving (250 0F for 20 min) and ethylene oxide gas (4 hours). The tested Niti wire was Sentalloy (GAC International).
Evaluation of sentalloy wire and dry heat sterilization demonstrated a significant increase in tensile strength when compared after 0, 1 and 5 cycles. Autoclaving santalloy wire also produced a significant increase in the tensile strength of the wire after 1 and 5 cycles. Ethylene oxide sterilization of sentalloy wires demonstrated no significant differences in tensile strength.
PROBLEMS ENCOUNTERED IN NICKEL TITANIUM ARCHWIRE /DISADVANTAGES
Expensive, particularly for newest products
Second highest arch wire-bracket friction after TMA
Difficult to place permanent bends and cannot bend wire over sharp edge or into complete loop
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Wires cannot be soldered and must be joined by mechanical crimping process.
Lowest in vitro corrosion resistance of wire alloys
Not self limiting – frequent visits necessary. Tendency of arch wire to slide from side to side, sometimes causing them to stick out beyond the terminal molar.
ADVANTAGES
Lowest force delivery of orthodontic wire alloys.
Excellent spring back in bending, particularly for super elastic and shape memory alloys.
Super elastic alloys can be heat treated by clinician to vary force delivery characteristics.
PROPRIETARY ARCH WIRES A)
A - COMPANY
1.
ALIGN TRUE
These are the only wires preformed in the true arch shape. Align is engineered with high elasticity to easily engage brackets, even on severely malposed teeth. It features exceptional shape memory and smooth low frictional surfaces. It also exerts continuous low forces. The arch wires are available in 0.14″ to 0.20″ round and 0.016″ x 0.016″ to 0.021″ x 0.025″ rectangular sizes, in small, medium and large arches.
2.
ALIGN NICKEL TITANIUM REVERSE CURVE OF SPEE ARCHWIRES
These wires are made with the same highly elastic material as the regular align nickel titanium wires. They provide the force needed to open the bite, close spaces
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and align the curve of spee. These wires are available in 0.016″ and 0.018″ round and 0.016″ x 0.022″ to 0.021″ x 0.025″ rectangular wires.
B)
AMERICAN ORTHODONTICS
1.
TITANIUM MEMORY WIRE
This is super elastic Nickel titanium wire available in several varieties. Standard memory arches feature natural arch form and two force levels. Force 1 (super elastic) from 0.016″ to 0.020″ round. Force 2 (high force) from 0.014″ to 0.018″ round.
Gold coated memory arches available only in 0.016″ round, are coated with 24 carat gold and are especially cosmetic with ceramic or plastic brackets.
Permanent centre line memory arches have a gable bend at the midline that acts as permanent reference point and a stop to prevent the wire from sliding through either central bracket during unscrambling. They come in 0.016″ and 0.018″ round sizes.
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Reverse curve of spee memory arches feature built in leveling and light continuous force for less wire bending and better control.
C)
DENTAURUM INTERNATIONAL
1.
REMITAN “LITE”
It is a super elastic nickel titanium wire with high elasticity and nearly continuous force over a wide deflection range. The smooth surface virtually eliminates friction and midlines are marked for easy placement. The wire delivers gentle force, making it particularly suited to leveling phase, can reduce treatment time and improve patient’s comfort. It is available in ideal arch form compatible with preadjusted appliances.
D)
GAC INTERNATIONAL
1.
NEO SENTALLOY
This wire provides the capability of 3 – dimensional control with a full size, single strand arch wire from the beginning of treatment. It offers outstanding shape memory and elasticity. Its heat activated light continuous forces are designed to produce the ideal biological tooth movement.
2.
SUPER BRAID
It is eight stranded braided super elastic Nickel titanium wire designed for initial 3- dimensional leveling of severely malposed teeth. It features low stiffness, predictable results, reduced chair time and patient’s comfort.
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E)
MASAL ORTHODONTICS INTERNATIONAL
1.
BENDABLE MASAL ALLOY (BMA)
BMA arches combine the continuous gentle force of titanium with the workability of steel. Elastic hooks, tear drop and bayonet bends can be bent into the wires, eliminating the need for auxiliaries.
BMA arch wires come in 0.016″ and 0.018″ round and 0.016″ x0.016″ to 0.019″ x 0.025″ rectangular sizes.
2.
ELASTINOL NICKEL TITANIUM ARCH WIRES
These wires retain their shape even when drastically deformed. They feature faster results, reduced chair time, low bracket friction and relatively low cost. They come in 0.012″ to 0.020″ round and 0.016″ x 0.016″ to 0.021″ x 0.025″ rectangular sizes.
3.
DRIFT FREE ELASTINOL
Drift free elastinol features a permanent 1 mm midline stop that reduces arch wire drifting and acts as a reference point. It comes in sizes of 0.016″, 0.018″ and 0.20″ round.
4.
ORTHOCOSMETIC ELASTINOL
These arches have an esthetic coating that blends well with ceramic or plastic brackets and resists staining, discoloration, cracking and chipping.
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5.
NITINOL
These arch wires are slightly stiffer than Elastinol, making them useful in the initial leveling stages. The natural shape is designed to reduce the need for contouring.
6.
RETROARCH
Retro arch reverse curve super elastic nickel titanium arch wires have a rocking chair shape that can open or close the bite quickly, consolidate the arch and eliminates the excess curve of spee. Toe in bends helps eliminate mesiolingual rotations and the extra wide form minimizes lingual drift of anterior teeth.
The arch wires are available in sizes of 0.014″ to 0.018″ round and 0.016″ x 0.016″ to 0.021″ x 0.025″ rectangular sizes.
F)
OREC CORPORATION
1.
NICKEL TITANIUM WIRES
Orec’s Nickel Titanium wire combines super elasticity with shape memory to provide optimum force distribution for leveling, aligning and rotation control.
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It enhances patient’s comfort and treatment efficiency while reducing bracket friction. Preformed upper and lower arches come in sizes from 0.014″ to 0.020″ round and 0.016″ x 0.022″ to 0.019″ x 0.025″ rectangular.
2.
SPEED ARCH WIRES
Nickel titanium and stainless steel speed arch wires have an arch form that reflects in-out effects of the speed appliance. The wires are designed to facilitate insertion into bracket and closure of speed bracket spring clips. Midlines are clearly marked and rounded edge is always directed towards the labial. Speed wires are available in upper and lower 0.017″ x0.022″ to 0.020″ x 0.025″ arch wire sizes.
G)
ORMCO CORPORATION
1.
NITI WIRE
Niti wire has an elastic range so great that it is virtually impossible to put a permanent set into the wire.
2.
REVERSE CURVE NITI
Reverse curve Niti has a shape that counters the extrusive component of space closing forces while continuing the bite opening process. It allows retraction without tipping of adjacent teeth into extraction sites or loss of incisor torque.
The posterior toe-in counteracts undesirable mesiolingual rotation of the molars. Extra archwidth prevents lingual collapse in the extraction sites.
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3.
NEW TURBO WIRES
This is a braided Nickel titanium arch wire that permits torque control from the first wire in treatment.
The braiding process increases the super elastic properties of Niti, so that a full size wire can be used in a very severe malocclusion without patient’s discomfort. The low stiffness of this wire makes it effective with ceramic as well as metallic brackets.
H)
ORTHO ORGANIZERS INTERNATIONAL
1.
NITANIUM ARCHES
Nitanium arches are Nickel titanium wires available in sizes from 0.014″ to 0.020″ round and 0.016″ x 0.022″ to 0.019″ x 0.025″ rectangular.
I)
ROCKY MOUNTAIN ORTHODONTICS
1.
ORTHONOL
Orthonol Nickel Titanium wire features great working range for fewer wire changes and adjustments, resistance to deformation, excellent shape memory and light continuous forces for patient’s comfort.
Its ultra smooth finish reduces bracket friction. Orthonol comes in 11 round and rectangular sizes of preformed arches and in 5 rectangular sizes of straight lengths.
J)
TP ORTHODONTICS
1.
REFLEX
Reflex super elastic Nickel Titanium arch wire is available in two arch forms.
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Curved memory arch (Reverse Curve of spee)
Straight arch (Natural)
Reflex is available in sizes from 0.014″ to0.020″ round and 0.016″ x 0.016″ to 0.021″ x0.025″ rectangular.
K)
UNITEK CORPORATION
1.
FLEXILOY
Flexiloy is made of a cobalt base Nickel alloy. In its initial work hardened temper, it is especially useful for making complex bends and loops. Heat treatment approximately doubles its spring temper. Flexiloy is available in two initial tempers, Blue and Yellow.
2.
NITINOL ACTIVE
Because of its unique formation it delivers the light, continuous forces needed for efficient tooth movement throughout treatment, yet with a slightly higher stiffness than most super elastic wires. The added control allows it to hold bends longer than other Nickel titanium wires.
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13
Beta Titanium Arch Wires
Beta titanium is the newest alloy to be introduced to orthodontic profession. Titanium has been used as structural metal since 1952, and its possible use in orthodontics has been suggested periodically.8,23,27,31
To compete with stainless steel, a wire must possess at least comparable formability and spring back, which is proportional to the ratio of yield strength to modulus of elasticity (Ys/E). 106
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The early industrial applications of titanium employed commercially pure material (99.2% titanium). At temperatures below 1,625 0 F this metal has a hexagonal closed packed (HCP) crystal form and an appliance constructed from pure titanium would have only 1/3rd the maximum elastic deflection of a comparable stainless steel appliance.
The second phase of titanium’s chronology saw the development of titanium alloys, but still based on HCP structure.
In the 1960,s an entirely different “high temperature” form of titanium alloy became available. At temperatures above 16250 F pure titanium rearranges into a body centered cubic (BCC) lattice, referred to as “Beta phase”. With the addition of elements such as molybdenum or columbium, a titanium based alloy can maintain its beta structure even when cooled to room temperature. Such alloys are referred to as beta stabilized titanium. The alloying and body centered cubic structure import a unique set of properties.
COMPOSITION
Titanium
77.8%
Molybdenum
11.3%
Zarconium
6.6%
Tin
4.3%
The alloy is marketed in the form of straight wire lengths or preformed arches under the trade name “TMA” or Titanium molybdenum alloy.
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MECHANICAL PROPERTIES
The mechanical properties of many titanium alloys can be altered by heat treatments that use the transformation from the α to β lattice structure. However, heat treatment of the current orthodontic β -titanium wire is not recommended.
Wrought Beta – titanium orthodontic wire has an elastic modulus of 71.7 Gpa and a yield strength between 860 and 1170 Mpa. These properties produce several clinically desirable characteristics. The low elastic modulus yields large deflections for low forces.
The high ratio of yield strength to elastic modulus produces orthodontic appliances that can sustain large elastic activations when compared with stainless steel devices of the same geometry.Beta – titanium can be highly cold – worked. The wrought wire can be bent into various orthodontic configurations and has formability comparable to that of austenitic stainless steel.
The modulus of elasticity of beta – titanium is approximately twice that of nitinol and less than one half that of stainless steel. Its stiffness makes its ideal in applications where less force than steel is required but where lower modulus materials would be inadequate to develop required force magnitudes.
Elasticity (10 MPa [GPa]
0.2% Offset yield strength (MPa)
Ultimate Tensile Strength
Number of 90 degree cold bends without fracture
S.S
179
1579
2117
5
Co-Cr
184
1413
1682
8
NiTi
41.4
427
1489
2
B-TITANUM
71.7
931
1276
4
Modulus of Alloys
3
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It has been shown that the formability of Beta titanium orthodontic wire, as measured by the ADA cold bend test, is similar to that of stainless steel. However the titanium alloy cannot be bent over as sharp a radius as stainless steel, so that some care in the selection of pliers and bending procedure is required. The Beta titanium wire can be joined by welding alone and has good corrosion resistance.
WELDING
Clinically, satisfactory joints can be made by electrical resistance welding of Beta titanium. Beta titanium joints of adequate strength and ductility can be produced with the standard commercial welders available to the orthodontist. Such joints need not be reinforced with solder.
A weld made with insufficient heat fails at the interface between the wires, whereas overheating may cause a failure adjacent to weld joint. Clinically a no of variables might cause joints of inconsistent strength with a given welder. Foremost among these factors are the condition of the electrodes, cleanliness of wire surfaces and proper positioning of the wires between the electrodes.
Flat to flat electronic configuration generally produces joints with considerably less distortion than is found with point to point arrangement. This electrode arrangement further stabilizes the wires as suggested by Burstone , and higher settings can be used with welders to obtain strong joints with less burning of metals.
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CLINCAL APPLICATION
Because of its unique and balanced properties, beta titanium wire can be used in a number of clinical applications. Ideal edge wise arches fabricated of titanium have significant superiority over stainless steel. They can be deflected approximately twice as far without permanent deformation, which allows a greater range of action for either initial tooth alignment or finishing arches. The forces which are produced are approximately .4 that of steel, producing a more gentle delivery of forces with an edge wise wire; for example an 0.018″ x 0.025″ wire in beta-titanium delivers about the same force as an 0.014″ x 0.020″ steel wire when activated in a second order direction. Furthermore, it would have the advantage of full bracket engagement and third order or torque control if used in a 0.018 ″ slot bracket. Beta titanium is ductile, which allows for placement of tie-back loops or complicated bends. Spring back properties are not lost during the bending operation and complicated configurations can be placed if needed.
The high ductility and formability of titanium allowed the placement of a vertical loop tie-back mesial to first molar as well as finishing bends with the arch.
The high ductility of Beta-titanium allows it to be formed into arches or segments with complicated loop configurations. A continuous arch with “T”, vertical, helical and “L” loops can be formed in small round wires.
In many applications, loop placement can better deliver the desired force system without side affects, than straight continuous wires. One of the advantages of beta-titanium, as used in loop configuration, lies with loop incorporation in larger 110
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cross-sections of edge wise wire which allows the loop to be positively oriented within the brackets.
Specialized springs or auxiliaries fabricated from beta-titanium allow for simplification in design in achieving identical force delivery. The low load deflection rate produced by the low modulus of elasticity and the high spring back allow a 12 mm activation to produce 60 gm of force in the midline without the placement of helices posteriorly, there by simplifying the design.
A high formability of titanium allows the fabrication of closing loops with or without helices. The low stiffness of the material and its high spring back improves a loop of any given design or allow for the maintenance of a given force system with simpler designs, as in elimination of helices or loops.
ION IMPLANTATION
A low coefficient of friction is usually desirable in an orthodontic arch wire. However studies have shown that TMA have a higher coefficient of friction than stainless steel. The friction is probably due to its relative softness compared to the harder stainless steel bracket. The surface treatment can increase the hardness and reduce the coefficient of friction of TMA wire while maintaining its desirable mechanical properties.
Ion Implantation is a process by which various elements or compounds are ionized and then accelerated towards a target, the orthodontic arch wire.
Ion implantation takes place in a vacuum chamber, where a vapour flux of ions is generated with an electron beam evaporator and deposited on the substrate. 111
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Gas ions (Nitrogen and oxygen) are simultaneously extracted from a plasma and accelerated in the growing physical vapour deposition film at energies of several hundred to thousand electrons volts.The ions penetrate the surface of the wire on impact, building up a structure that consists of both the original wire and a layer of tin compounds on the surface and immediate subsurface. This la yer is extremely hard and creates considerable amount of compressive forces in the material at the atomic level. The compressive forces and increased surface hardness improves the fatigue resistance and ductility and reduce the coefficient of friction of the wire. The superficial compressive forces also minimize any effects of surface flaws. Implantation produces no sharp interface between coating and wire which can lead to bond failure and it does not alter wire dimensions.
Implantation can take place at relatively low temperatures from subzero to 7000C which allows improvement of surface characteristics without degradation of other mechanical properties. The thickness of the implanted surface can be precisely controlled.
Two varieties of TMA-low friction and colored wires were produced by varying the type and thickness of ions. Studies have shown that surface treatment by ion implantation can maintain all the desirable properties of TMA and can improve its ductility and its resistance to fracture and fatigue. At the same time it reduces high coefficient of friction to about the same level as that of stainless steel.
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14
Chinese Ni-Ti Wire A new nickel titanium alloy has been developed especially for orthodontic applications by Dr. TIEN HUA CHENG and associates at the General Research institute for non ferrous metals in Beijing, China.
This alloy has unique characteristics and offers significant potential in the design of orthodontic appliances. Its history of little work hardening and a parent phase, which is austenite, yield mechanical properties that differ significantly from nitinol wire. In addition, Chinese niti wire has a much lower transition temperature than nitinol wire.9 This wire is also called as 27 0 C super elastic copper NiTi. It contains alloy additions of nominally 5 to 6 % copper and .2 to .5 % chromium. According to its manufacturer, this product is an austenitic active wire whose copper addition increases its strength. Unfortunately this occurs at the expense of increasing its phase transformation temperature above that of the oral cavity. To compensate for this
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unwanted effect .5 % chromium is added to return the transformation temperature to 270C.
Two other alloys are also available from this family of Nickel – Titanium – Copper – Chromium alloys. One that has a transformation temperature of 35 0C and other that contains .2 % chromium and transforms at 40 0C.
Because the transformation temperature of these latter two wires are higher than the before mentioned first wire, they will increasingly be influenced by temperature as they represent the thermoelastic Nitinol described before.
MECHANICAL PROPERTIES
1)
The wire has a spring back that is 4.4 times that of comparable stainless steel wire and 1.6 times that of nitinol wire, if spring back is measured at yield based on a 5 mm span cantilever test.
2)
At 80% of activation, the average stiffness of Chinese Niti wire is 73% that of stainless steel wire and 36% that of nitinol wire.
3)
Unlike wires of other orthodontic alloys, the characteristic stiffness is determined by the amount of activation. The load deformation rate at small activations is considerably higher than that at large activations. 0
The stiffness is approximately the same between room temperature at 22 C and mouth temperature at 37 0C.
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4)
Chinese Niti wire deformation is not particularly time dependent and unlike Nitinol wire, will not continue to deform a significant amount in the mouth between adjustments.
5)
Chinese Niti wire is highly suitable if low stiffness is required and large deflections are needed.
Its higher stiffness at small activations make it more effective than wires of traditional alloys whose force levels may be too low (as teeth approach the passive shape of the wire).
CLINICAL SIGNIFICANCE
Because of its high range of action or spring back, Chinese niti wire is applicable in situations where large deflections are required. Applications include straight wire procedures when teeth are badly malaligned and in appliances designed to deliver constant forces during major stages of tooth movements. The amount of deformation without notable permanent set is remarkable – 4.4 times that of stainless steel wire and 1.6 times that of nitinol wire.
Achievement of relatively constant forces has been obtained traditionally by lowering the load deflection rate of the orthodontic appliance. This has been accomplished by configurational design; for instance, placing helices or additional wire in the appliance.
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15
J apanese Ni-Ti Alloy Arch Wires In 1978, Furukawa Electric company limited of Japan produced a new type of Japanese NiTi alloy possessing excellent spring back, shape memory and super elasticity.
MECHANICAL PROPERTIES
The Japanese niti alloy wire has higher values of elastic modulus than the nitinol wire. When the stretch exceeds 2 %, the stress value does not change appreciably. When the strain was induced at 8%, it produces stresses of 55 to 58 kg/mm2. When the wire specimen was then stretched for more than 8%, the stress was increased further. This property is called as super elasticity.
When strain was reduced, the stainless steel, Co-Cr-Ni and nitinol wires all exhibit almost straight stress strain curves. In comparison when strain was reduced, the Japanese Niti alloy wire did not changes proportionally to the stress decrease from 8% to 2%. There was no permanent set when the stress reached zero.
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Heat treatment of Japanese niti alloy does make a dramatic change in its mechanical property. To attain optimal use of super elastic property in clinical orthodontics, the influence of various series of heat treatment was studied. When the heat application was raised to 500 0C, the force level indicating the super elastic property can be reduced.
Thus, arch wires providing a different magnitude of force can be fabricated from the wires of same diameter. In addition, in the preformed arch wire, different magnitude of force can be produced by controlling the temperature and time in the desired section of arch wire.
Japanese niti possesses three good mechanical properties;
excellent spring back
shape memory
super elasticity
CLINICAL APPLICATION
Since the metallurgical tests have determined that Japanese niti alloy wire is potentially useful and effective in clinical orthodontics, setting arch wire have been fabricated to enhance the efficiency of multi bracketed technique.
By evaluating clinical experience with the Japanese niti alloy wire, many possibilities exist with the use of its super elastic property.
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16
Alpha Titanium Alloy Arch Wires It is the recent alloy in the family of titanium alloys. Its composition is: Titanium
90%
Aluminum
6%
Vanadium
4%
The alloy is different in that its molecular structure resembles a closely packed hexagonal lattice as against the BCC lattice of TMA. The hexagonal lattice possess fewer slip planes. Slip planes are clusters of atoms in a crystal that glides past one another during deformation. More the slip planes, the easier it is to deform the material. BCC structures are defined as having two slip planes where hexagonal lattice has only one active slip plane along its base. Thus the near alpha phase titanium alloy is less ductile than TMA. The alloy is strictly near Alpha phase titanium alloy rather than a pure alpha titanium because there is a certain amount of Beta phase retained in them at room temperature. 118
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17
Copper Ni-Ti Alloy Arch Wires Copper Niti was introduced by Rohit Sachdeva and Suchio Mriyasaki in 1994.
It’s a new quaternary alloy (Nickel, Titanium, Copper and Chromium) with different advantages over the formerly available Nickel Tit anium alloys.
Copper niti is more resistant to permanent deformation compared with other Nickel-Titanium alloys. It exhibits better spring back characteristics. It exhibits a smaller drop in tooth driving force than other Nickel titanium alloys. It generates a more constant force over long activation spans than other Nickel titanium alloys and does so on a consistent basis. Addition of copper combined with more sophisticated manufacturing and thermal processes make possible the fabrication of four different copper niti arch wires with precise and consistent transformation temperatures
150C, 270C , 350C and
400C. This enables the clinician to select arch wires on a case specific basis.
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COMPOSITION
Titanium
43%
Nickel
49.86%
Chromium
.5%
Copper
5.64%
Copper niti delivers more constant forces especially for small activations compared to super elastic wires. It makes possible the insertion of larger size wires, and better bracket slot engagement early in treatment without causing pain and discomfort.
The surface of cu-niti is quite porous and rough. It resembles the surface of untreated TMA wire.
Depending upon transformation temperatures / austenitic finish temperature cu-niti can be classified into
Type I austenitic finish
150C
Type II
,,
270C
Type III
,,
350C
Type IV
,,
400C
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VARIABLE TRANSFORMATION TEMPERATURE
The stability of the martensite and austenitc phase at a given temperature is based upon transformation temperature of the alloy. One of the most important markers is the materials austenitic finish temperature.(Af)
It is the difference between Af temperature and mouth temperature that determines the force generated by Nickel Titanium alloys. Af temperature can be controlled over a wider range by affecting the composition, thermo mechanical treatment and manufacturing process of the alloy.
This alloy has the advantage of generating more constant forces than any other super elastic Nickel Titanium alloy.
It is more resistant to deformation as a result of thermo mechanical insults in the mouth.
Type I is not used for clinical applications due to high force level. Type II produces the optimum force and is indicated in normal patients. Type III is indicated in patients with a low to normal threshold of pain and also in periodontically compromised patients. Type IV produces the lowest level of force and are good in patients who are highly sensitive to pain and periodontically compromised compromised patients.
Quick and simple trick is to apply ice to the section of arch wire and can be placed into the bracket easily.
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Cu-Niti is supplied in various sizes. 270 C 0.014″ 0.014″, 0.016″ 0.016 ″, 0.018″ 0.018 ″,0.016″ ,0.016 ″, 0.022″ 0.022 ″, 0.017″ 0.017 ″ x 0.025″ 0.025 ″, 0.017″ 0.017 ″ x 0.025″ 0.025 ″. 350 C 0.016″ 0.016″, 0.018″ 0.018 ″, 0.016″ 0.016 ″ x 0.022″ 0.022 ″, 0.017″ 0.017 ″ x 0.022″ 0.022 ″, 0.017″ 0.017 ″ x 0.025″ 0.025″. 400 C 0.016″ 0.016″ x 0.022″ 0.022″, 0.017″ 0.017″ x 0.025″ 0.025 ″, 0.019″ 0.019″ x 0.025″ 0.025″.
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18
O ptiflex Arch Wire Optiflex is a recently introduced arch wire by Tallas. It combines highly
aesthetic appearance with unique mechanical properties.
It is made of clean optical fiber and consisof 3 layers.
Silicon-di-oxide core that provides the force for moving teeth.
Silicon resin middle layer that protects the core from moisture and adds strength.
Strain resistant nylon outer layer that prevents damage to the wire and further increases its strength.
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The wire is manufactured in various sizes, and can be either round or rectangular. Sharp bends must be avoided, since they could fracture the core. It is highly resilient arch wire that is especially effective in the alignment of crowded teeth. It has got a wide range of action and apply light continuous force. Lee white arch wire of Lee pharmaceuticals is tooth colored epoxy coated arch wire that has superior wear resistance.
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19
Dead Soft Security Arch Wires It has been introduced recently by Binder and Scott . In a non-extraction case, an arch wire is usually placed to initiate tooth movement immediately after bonding. However,in an extraction case a proper arch wire might create undesired tooth movement before extractions are performed. This problem can be avoided by placing sectional arches made of dead soft brass wire or twisted double strands of 0.008″ or 0.010″ dead soft stainless steel ligature wires. These arches are bend to lie passively in all attachments.
The same type of sectional arches can be used as final arch wires in one or both arches in conjugation with snake elastics to enhance intercuspation prior to appliance removal.
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20
Conclusion Recent advances in orthodontic arch wires proves a clear commitment to high performance standard, lifelong learning and strict accreditation in search of an ideal wire. To obtain benefit of optimum and predictable treatment results one can depend on selection of appropriate wire size and alloy type.
The eminent orthodontic campaigners have come so far in search of an ideal wire with which they can play with, instead, it playing with them and perhaps someday these wire benders will come up with long overdue research of an ideal wire and we wait that the dawn will come early.
Success is not final, failure is not fatal: it is the courage to continue that counts.
”
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