Сraig. Dental Materials

Impression materials

Alginate

Polyether

Polysulfide

Silicone

Maxillofacial materials

Polyurethane

Polyvinylchloride

Silicone rubber

*Crosshead speed, 2 cm/min

separate phases. The trend in the development of materials for various applications is toward composites rather than completely new classes of materials. There can be metal, ceramic, and polymer-based composites. Important examples of dental composites include posterior resin composite used as direct esthetic restorative materials. Such composites are made from an organic polymer matrix (usually a diacrylate) filled with an inorganic phase, such as borosilicate or strontium glass, lithium or barium, aluminum silicate, or colloidal silica.

Factors that affect the properties of composites include: (1) the state of matter of the second (dispersed) phase; (2) the geometry of the second phase; (3) the orientation of the second phase; (4) the composition of the dispersed and continuous phases; (5) the ratio of the phases; and (6) bonding of the phases. Examples of properties that can be changed (improved if the composites are judiciously developed) are:

(1) modulus, (2) strength, (3) fracture toughness,

(4) wear resistance, (5) thermal expansion, and

(6) chemical and corrosion resistance.

A simple example of how adding a second phase affects properties is now illustrated. Consider a series of continuous parallel glass fibers oriented in the same direction in a plastic matrix. If a tensile load is applied to the specimen in the

direction of the fibers, the elastic modulus of the composite E, is

where E, and Em and V, and Vm represent the elastic modulus and volume fraction of the fiber and matrix. If, on the other hand, the tensile load is applied in the direction transverse to the fibers, the composite elastic modulus would be

If the volume fraction of fibers is zero (i.e., the material is strictly a polymer), the modulus is that of the polymer, and if the volume fraction is loo%, the material is a glass and has the modulus of glass. Thus the moduli of the polymer and glass serve as lower and upper bounds on the composite modulus. Furthermore, from the above two equations, we see that in addition to the ratio of the two phases, the orientation of the second phase also plays an important role in the composite properties.

The basis for the function of a dispersed phase in a matrix is shown in Fig. 4-31. A single fiber is shown surrounded by a matrix, and the tensile stress in the fiber is plotted versus the distance along the fiber. The load is applied to the matrix and is transferred to the fiber by shear at the interface. The elastic and plastic deformation in the matrix can be transferred to the fiber if the modulus of the matrix is less than that of the fiber. Also, the bond between the matrix and the fiber must be maintained or the stress will drop to the frictional force. As the load is increased, the tensile stress in the fiber may reach the ultimate shear stress and the fiber will fail.

Continually increasing the volume fraction of the fibers should continue to increase the strength of the composite. However, as the concentration of fibers increases, more and more contact of the fibers with each other occurs, and

Fig. 4-31 Stress on a glass fiber in a plastic matrix.

(Adapted from Titelman AS, McEvily AJ,Jr: Fracture of structural materials, New York, 1967, John Wiley & Sons,

pp 635-661.)

premature rupture results. Therefore for many composites, the maximum strength occurs at a volume fraction of 80% for the dispersed phase.

As a further illustration of the factors that affect the properties of a composite, consider the filled resins used in dentistry. For many of these dental composites a random arrangement of the dispersed phase is used, even though a random orientation results in about a sixfold lower strength compared to an oriented dispersed phase. However, the resultant lower strength due to random second phase orientation can be counteracted in several ways. The use of a fine (e.g.,<1ym) dispersed phase increases strength. Also, the selection of the shape of the dispersed particles is important, with rods and plates being more effective than spheres in improving the strength. The principal factors needed are (1) a high-strength dispersed phase; (2) a more ductile matrix phase; (3) fine dispersed particles at the optimum volume fraction; and (4) adhesion between the dispersed and matrix phases. The last requirement is usually accomplished by treating the dispersed phase with an organosilane. The silanes, often called coupling agents, react with the glass or water adsorbed on the glass and form a bond with the resin.

In our discussion so far, we have introduced and discussed mechanical properties that are mainly dependent on the bulk characteristics of a material. In this section, mechanical properties that are more a function of the surface condition of a material are presented. In particular, the concepts of hardness, friction, and wear are summarized.

HARDNESS

The property of hardness is one of major importance in the comparison of restorative materials. Hardness may be broadly defined as the resistance to permanent surface indentation or penetration.

Formulating a more rigorous definition of hardness is difficult because any test method will, at a microscopic level, involve complex surface morphologies and stresses in the test material, thereby involving a variety of qualities in any single hardness test. Despite this condition, the most common concept of hard and soft substances is their relative resistance to indentation. Hardness is therefore a measure of the resistance to plastic deformation and is measured as a force per unit area of indentation (Fig. 4-32).

Based on this definition of hardness, it is clear why this property is so important to dentistry. Hardness is indicative of the ease of finishing of a structure and its resistance to in-service scratching. Finishing or polishing a structure is important for esthetic purposes and, as discussed previously, scratches can compromise fatigue strength and lead to premature failure.

Some of the most common methods of testing the hardness of restorative materials are the Brinell, Knoop, Vickers, Rockwell, Barcol, and Shore A hardness tests. Each of these tests differs slightly from the others, and each presents certain advantages and disadvantages. They have a common quality, however, in that each depends on the penetration of some small, symmetrically shaped indenter into the surface of the material being tested. The various hardness tests differ in the indenter material, geometry, and load. The indenter may be made of steel, tungsten carbide or diamond, and be shaped as a

sphere, cone, pyramid, or needle. Loads typically range from 1to 3000 kg. The choice of a hardness test depends on the material of interest, the expected hardness range, and the desired degree of localization.

The general procedure for testing hardness, independent of the specific test, is as follows. A standardized force or weight is applied to the penetrating point. Applying such a force to the indenter produces a symmetrically shaped indentation, which can be measured under a microscope for depth, area, or width of the indentation produced. The indentation dimensions are then related to tabulated hardness values. With a fixed load applied to a standardized indenter, the dimensions of the indentation vary inversely with the resistance to penetration of the material tested. Thus lighter loads are needed for softer materials.

Brinell Hardness Test The Brine11 hardness test is among the oldest methods used to test metals and alloys used in dentistry. The method depends on the resistance to the penetration of a

small steel or tungsten carbide ball, typically 1.6 mm in diameter, when subjected to a weight of 123 N. In testing the Brinell hardness of a material, the penetrator remains in contact with the specimen tested for a fixed time of 30 seconds, after which it is removed and the indentation diameter is carefully measured. A diagram showing the principle of Brinell hardness testing, together with a microscopic view of the indentations into a gold alloy, is shown in Fig. 4-33. The resulting hardness value, known as the Bri- nell hardness number (BHN), is computed as a ratio of the load applied to the area of the indentation produced. The formula for computing the BHN is as follows:

L

BHN = nD

2

In this formula L is the load in kilograms, D is the diameter of the ball in millimeters, and d is the diameter of the indentation in millimeters, thus the units for BHN are kg/mm2. The smaller

Condensed gold

Foil

Powdered

Gold alloys*

Type 1

Type I1

Type I11

Type IV

40% Au-Ag-CU

99% noble alloyt

*Alloysthat are hardenable by heat treatment are in the hard condition.

tFor metal-ceramic restorations.

the area of indentation, the harder the material and the larger the BHN value. Tables of Brinell hardness values have been developed from this formula for indentations of different diameters. Because the Brinell hardness test yields a relatively large indentation area, this test is good for determining average hardness values and poor for determining very localized values. The Brinell hardnesses of some dental casting alloys and condensed gold are listed in Table 4-15.

The Knoop hardness test was developed to fulfill the needs of a microindentation test method. A load is applied to a carefully prepared diamond indenting tool with a pyramid shape, and the lengths of the diagonals of the resulting indentation in the material are measured. The shape of the indenter and the resulting indentation are illustrated in Fig. 4-34, A . The Knoop hardness number (KHN) is the ratio of the load applied to the area of

the indentation calculated from the following formula:

L

KHN = -

12c,

In this equation L is the load applied, 1 is the length of the long diagonal of the indentation, and C, is a constant relating 1 to the projected area of the indentation. The units for KHN are also kg/mm2. Similar to the Brinell method, higher values for KHN represent harder materials.

The Knoop method is designed so varying loads may be applied to the indenting instrument. The resulting indentation area, therefore, varies according to the load applied and the nature of the material tested. The advantage of this method is that materials with a great range of hardness can be tested simply by varying the test load. Because very light load applications produce extremely delicate microindentations, this method of testing can be employed to examine materials that vary in hardness over an area of interest. For example, the Knoop method has been used extensively in testing the hardness of both enamel and dentin in extracted teeth and in determining the hardness of metals and alloys that have isolated hard and soft phases throughout the material. The chief disadvantages of the method are the need for a highly polished and flat test specimen and the time required to complete the test operation, which is considerably

Fig. 4-34 A, Principle of the Knoop hardness measurement; 0, the diamond pyramid (Vickers) indentation test.

Silicon carbide abrasive Feldspathic porcelain Cobalt-chromium partial denture

alloy Enamel Gold foil Dentin Cementum

Zinc phosphate cement Denture acrylic

greater than that required for some other, less precisely controlled methods. The KHNs of some dental materials are listed in Table 4-16.

Vickers Hardness Test The Vickers hardness test, or 136-degree diamond pyramid, is also suitable for testing the surface hardness of materials. It has been used to a limited degree to test the hardness of restorative dental materials. The method is similar in principle to the Knoop and Brinell tests, except that a 136degree diamond pyramid-shaped indenter is forced into the material with a definite load application. The indenter produces a square in-

I06 Chapter 4 MECHANICAL PROPERTIES

dentation, the diagonals of which are measured as shown in Fig. 4-34, B. Equipment for Knoop hardness testing has been adapted to use the 136-degree indenter. Loads are varied from 1 to 120 kg, depending on the hardness of the tested material. The Vickers test is especially useful in measuring the hardness of small areas and for very hard materials.

Rockwell Hardness Test The Rockwell hardness test was developed as a rapid method for hardness determinations. A ball or metal cone indenter is normally used, and the depth of the indentation is measured with a sensitive dial micrometer. The indenter balls or cones are of several different diameters, as well as different load applications (60 to 150 kg), with each combination described as a special Rockwell scale, Rockwell A-G, denoted R,, R,, etc.

The superficial Rockwell method has been used to test plastics used in dentistry. This method uses a relatively light (30 kg) load and a large-diameter (12.7 mm) ball in comparison with the standard Rockwell methods. The test is made by first applying a preload (minor load) of 3 kg. A major load of 30 kg then is applied to the specimen for 10 minutes before a reading is taken. Because dental plastics are viscoelastic, recovery of the indentation occurs once the major load has been removed. The percent recovery can be determined on the same specimen by the following equation:

A - B

Percent recovery = -X 100%

A

where A is the depth of the indentation caused by application of the major load for 10 minutes, and B is the depth of the indentation after the major load has been removed for 10 minutes. Values of indentation depth and percent recovery for some dental plastics are listed in Table 4-17. The advantages of the Rockwell hardness test are that hardness is read directly and it is good for testing viscoelastic materials. The disadvantages are that a preload is needed, greater time is required, and

the indentation may disappear immediately on removal of the load.

Barcol Hardness Test Barcol hardness is one method used to study the depth of cure of resin composites. The Barcol indenter is a springloaded needle with a diameter of 1 mm that is pressed against the surface to be tested. If no penetration of the needle into the surface occurs, the scale reads 100. The reading on the scale decreases as the indenter penetrates the surface. Depth of cure of a resin composite is tested by preparing specimens varying in thickness from 0.5 to 6.0 mm or more in increments of 0.5 mm. Then the top surface of a specimen is activated by a light-curing unit. The Barcol hardness of the top surface is compared with that of the bottom surface. The depth of cure is defined as the maximum thickness at which the Barcol reading of the bottom surface does not change by more than 10% of the reading of the top surface. Research has shown that a 10% decrease in Barcol hardness of a resin composite results in a 20% decrease in the flexural strength.

Shore A Hardness Test The hardness measurements described previously cannot be used to determine the hardness of rubbers, because the indentation disappears after the removal of the load. An instrument called a Shore A Durometer is used in the rubber industry to

determine the relative hardness of elastomers. The instrument consists of a blunt-pointed indenter 0.8 mm in diameter that tapers to a cylinder 1.6 mm. The indenter is attached by a lever to a scale that is graduated from 0 to 100 units. If the indenter completely penetrates the specimen, a reading of 0 is obtained, and if no penetration occurs, a reading of 100 units results. Because rubber is viscoelastic, an accurate reading is difficult to obtain because the indenter continues to penetrate the rubber as a function of time. The usual method is to press down firmly and quickly on the indenter and record the maximum reading as the Shore A hardness. The test has been used to evaluate soft denture liners, mouth protectors, and maxillofacial elastomers, values of which are listed in Table 4-18.

NANO-INDENTATION

Traditional indentation tests use loads as high as several kilograms and result in indentations as large as 100 pm. Although valuable for screening materials and determining relative values among different materials, these tests are subject to limitations. Many materials have microstructural constituents or, in the case of microfilled composites, filler phases substantially smaller than the dimensions of the indenter. To accurately measure the properties of these microphases, it is necessary to be able to create indentations of a smaller size scale and to spatially control the location of the indentations. In this regard, special indenta-

tion techniques have recently been introduced. These techniques, commonly referred to as nanoindentation, are able to apply loads in the range of 0.1 to 5000 mg, resulting in indentations approximately 1 ym in size. In addition, indentation depth is continuously monitored, obviating the need to image the indentation to compute mechanical properties. Although most commonly used to measure hardness of micron-sized phases, the technique is also useful for measuring modulus. For brittle materials, yield strength and fracture toughness may be determined.

The nano-hardness, dynamic hardness, and elastic moduli values of human enamel and dentin are listed in Table 4-19, along with the nanohardness and elastic modulus for the region of the dentin-enamel junction. The nano-hardness of dentin of 71 kg/mm2 agrees well with the Knoop value of 68 kg/mm2 reported in Table 4-16;however, the nano-hardness of 457 kg/mm2 for enamel is considerably higher than the Knoop value of 343 kg/mm2. This difference may result from the much smaller indentation used in the nano-indentation test in relation to the size of the enamel rods. The dynamic hardness values are lower than those for the corresponding nanohardness because they are calculated from the maximum displacement, whereas the nanohardness values are calculated from the permanent deformation. The elastic moduli of 87.7 and 24.0 GPa for enamel and dentin by nanoindentation are in reasonable agreement with the values from compressive test specimens of 84.1 and 18.3 GPa, respectively. Of special interest is the elastic modulus of 53.2 GPa for the region of the dentin-enamel junction, which is intermediate to the values for enamel and dentin. The nano-indentation test is especially useful in studying this small region, which was not possible with compressive or tensile tests.

FRICTION

Friction is the resistance to motion of one material body over another. If an attempt is made to move one body over the surface of another, a restraining force to resist motion is produced

(Fig. 4-35). This restraining force is the (static) frictional force and results from the molecules of the two objects bonding where their surfaces are in close contact. The frictional force, F,, is proportional to the normal force (F,) between the surfaces and the (static) coefficient of friction (~1,).

The coefficient of friction varies between 0 and 1 and is a function of the two materials in contact, their composition, surface finish, and lubrication. Similar materials have a greater coefficient of friction, and if a lubricating medium exists at the interface, the coefficient of friction is reduced.

The conditions for motion are for the applied force to be greater than F,. Once motion occurs, molecular bonds are made and broken, and microscopic pieces break off from the surfaces. With motion, a sliding or kinetic friction is produced, and the force of kinetic friction opposes the motion.

Frictional behavior therefore arises from surfaces that, because of microroughness, have a small real contact area (see Fig. 4-35). These small surface areas result in high contact stresses, which lead to local yielding. The resistance to shear failure of the junctions results in the fric-

Fig. 4-35 Microscopic area of contact between two objects. The frictional force, which resists motion, is proportional to the normal force and the coefficient of friction.

(Adapted from Tipler PA: Physics, 1976, Worth, p 156.)

tional force. When static friction is overcome and relative motion takes place, it is accompanied by the modification of the interface through kinetic friction and wear.

An example of the importance of friction in dentistry lies in the concept of roughening the surface of a dental implant to reduce motion between the implant and adjacent tissue. It is perceived that a rough surface and resultant less

motion will provide for better osseointegration. Friction is also important in sliding mechanics used in the orthodontic movement of teeth.

WEAR

Wear is a loss of material resulting from removal and relocation of materials through the contact of two or more materials. When two solid materials are in contact, they touch only at the tips of their highest asperities (see Fig. 4-35). Wear is usually undesirable, but under controlled conditions during finishing and polishing procedures, wear is highly beneficial.

Several factors make wear of biomaterials unique. Most importantly, wear can produce biologically active particles, which can excite an inflammatory response. The wear process can also produce shape changes that can affect function. For example, wear in the oral cavity is characterized by the loss of the original anatomical form of the material. Wear of tooth structure and restorative materials may result from mechanical, physiological, or pathological conditions. Normal mastication may cause attrition of tooth structure or materials, particularly in populations that consume unprocessed foods. Bruxism is an example of a pathological form of wear in which opposing surfaces slide against each other. If improperly performed, tooth brushing with a dentifrice may cause an abrasive form of wear.

Wear is a function of a number of material and environmental factors, including the nature of wearing surfaces (i.e., inhomogeneity, crystal orientation, phases, and inclusions present); the microscopic contact; interaction between sliding surfaces (i.e., elevated stress, temperature, and flow at contact points, leading to localized yielding, melting, and hardening); lubrication; and different material combinations. In general, wear is a function of opposing materials and the interface between them. The presence of a lubricating film, such as saliva, separates surfaces during relative motion and reduces frictional forces and wear.

In general, there are four types of wear: (1) ad-

hesive wear; (2) corrosive wear; (3) surface fatigue wear; and (4) abrasive wear. Adhesive wear is characterized by the formation and disruption of microjunctions. The volume of wear debris, 1.: is determined by

where xis the total sliding distance, k is the wear coefficient, FL is the perpendicular force, a n d p is the surface hardness (of the softer material).

Abrasive wear involves a soft surface in contact with a harder surface. In this type of wear, particles are pulled off of one surface and adhere to the other during sliding. There can be two types of abrasive wear: twoand three-body abrasion (wear). This type of wear can be minimized if surfaces are smooth and hard and if particles are kept off the surfaces. Corrosive wear is secondary to physical removal of a protective layer and is therefore related to the chemical activity of the wear surfaces. The sliding action of the surfaces removes any surface barriers and causes accelerated corrosion. In surface fatigue wear, stresses are produced by asperities or free particles, leading to the formation of surface or subsurface cracks. Particles break off under cyclic loading and sliding.

In general, metals are susceptible to adhesive, corrosive and three-body wear, whereas polymers are susceptible to abrasive and fatigue wear.

Wear has been studied by (1) service or clinical testing, (2) simulated service measurements,

(3)model systems using various wear machines,

(4)measurements of related mechanical properties such as hardness, and (5) examination of the amount and type of surface failure from a single or low number of sliding strokes.

Two-body abrasion tests have been used to rank the wear resistance of restorative materials. As shown in Table 4-20, the resistance of composite resins to abrasion depends on the nature of the filler particles (glass or quartz) and on silanation of the filler. Three-body abrasion tests

are often used to compare the abrasion resistance of tooth structure with that of dentifrices and

prophylaxis materials. Enamel is about 5 to 20 times more resistant to abrasion than dentin. Cementum is the least resistant to abrasion. Measurements of enamel loss during a 30-second prophylaxis have shown that fluoride is removed from the enamel surface and have allowed estimation of the removal of enamel to be 0.6 to 4 pm, depending on the abrasive.

Unfortunately, a 1:l ratio between wear observed clinically and that measured in the laboratory seldom exists. Thus most tests strive to rank materials in an order that is seen clinically. Traditional wear tests measure the volume of material lost but do not reveal mechanisms of wear, whereas a single-pass sliding technique may characterize modes of surface failure. In general, wear data do not correlate well with other mechanical property data, making it more difficult to infer wear properties from other, simpler, laboratory tests.

In the study and evaluation of wear, note that multiple processes occur simultaneously, and materials, mechanics, and environment have combined effects on wear. Most important is the fate of the wear particles. Are these particles

The mechanical properties of dental restoration materials must be able to withstand the stresses and strains caused by the repetitive forces of mastication. The design of dental restorations is particularly important if the best advantage of a material is to be taken. The necessary designs are those that do not result in stresses or strains that exceed the strength properties of a material under clinical conditions.

Stresses in dental structures have been studied by such techniques as brittle coatings analysis, strain gauges, holography, twoand threedimensional photoelasticity, finite element analysis, and other numerical methods. Stress analysis studies of inlays, crowns, bases supporting restorations, fixed bridges, complete dentures, partial dentures, endodontic posts, and implants have been reported, as well as studies of teeth, bone, and oral soft tissues. The stress analysis literature is too extensive and beyond the scope of this text to review. Only brief summaries of two-dimensional photoelasticity and finite element analysis, and the advantages and disadvantages of both, are provided.

TWO-DIMENSIONAL PHOTOELASTICITY

The procedure for two-dimensional models is to prepare a transparent plastic or other isotropic model of the restoration or appliance. The material becomes anisotropic (its properties exhibit a directional dependence) when stressed, so the behavior of light is affected by the direction it takes. As a result of the applied stress, the plastic model exhibits double refraction because of its anisotropic structure. The light from a source passes through a polarizer, which transmits light waves parallel to the polarizing axis, known as plane polarized light. The plane polarized light is converted to circularly polarized light by a