Biomaterials for Dental Applications

Abstract

Based on population projections, the number of teeth at risk for dental disease in 1980 was 2.8 billion for those over the age of 65, and is expected to increase to 5 billion teeth at risk by 2020 (http://www.ncbi.nlm.nih.gov). Dental materials are going to play an important role in the treatment of such ailments. As a matter of fact, dental materials have a long checkered history. From a historical viewpoint, many different materials, both natural and artificial, have been used in treating various oral disorders. Obviously for the materials to succeed, they have to be at least biocompatible and should be able to bond with teeth or bone. These materials should appear like natural tooth, if visible, and exhibit properties similar to those from various parts of a tooth.

14.1 Introduction

Based on population projections, the number of teeth at risk for dental disease in 1980 was 2.8 billion for those over the age of 65, and is expected to increase to 5 billion teeth at risk by 2020 [1]. Dental materials are going to play an important role in the treatment of such ailments. As a matter of fact, dental materials have a long checkered history. From a historical viewpoint, many different materials, both natural and artificial, have been used in treating various oral disorders. Obviously for the materials to succeed, they have to be at least biocompatible and should be able to bond with teeth or bone. These materials should appear like natural tooth, if visible, and exhibit properties similar to those from various parts of a tooth.

While dental materials are, in general, classified as preventive materials or restorative materials, the primary use of dental materials is for restoration. Dental restorative materials are made up of different types of materials such as metals, ceramics, polymers, and composites. The selection of the type of dental restorative material is dependent on many factors, such as the patient, the dentist, the material, and characteristics of the tooth itself. Important parameters for choosing dental restorative materials include strength, hardness, wear resistance, chemical resistance, and esthetic appearance. Metals are primarily used where strength and durability are needed. Typical uses of metals include amalgams, casting alloys, and porcelain fused to metal (PFM) restorations. Ceramics are also used as in conjunction with metal as PFM or as stand-alone all-ceramic materials. They are used where chemical durability and esthetics are important. They are used as veneers, crowns, protective layers, etc. Polymeric materials are used as cements and as cavity fillings in conjunction with particulate ceramic filler materials. Polymers are useful because of esthetics and adhesion properties.

This chapter will review important features of metals, ceramics, polymers, and composites in restorative dentistry. The subsections will describe the compositions, the relevant materials, scientific aspects, and related materials properties. Where appropriate, specialty processing techniques associated with manufacturing/fabrication will be described.

14.2 Historical Perspectives

The use of both artificial and natural materials in dentistry is well known for about 5000 years [2]. Around 700 B.C. Etruscans used elephant or hippopotamus ivory to construct partial dentures fixed to natural teeth by gold wires. Cavities in teeth were filled with a range of materials such as gum, metals, ivory, etc. Fossil evidence indicates that 2000 years ago, the art of dental implantology was already taking off by replacing lost teeth with alloplastic or homologous materials such as human and animal teeth, carved human and animal. These replacements were not only functional but they were used to meet esthetic needs. Ancient dental implant designs emerged in the Middle East, Western Europe, Asia, and Central and South America [3].

About 1000 years ago, the Spaniard Alabucasim recommended the replacement of teeth by transportation (transfer of living tissue), indicating that this was an acceptable strategy for replacing missing teeth. Several hundred years ago in France and England, it was fashionable to replace lost teeth with transplanted teeth obtained from young people who were paid for their extracted teeth [4]. In 1774, Duchateau and de Chemont developed a process for making porcelain dentures. In 1808, Fonzi, an Italian dentist, produced an individual porcelain tooth. In 1816, Taveau developed the first amalgam in France. A widespread use of vulcanized rubber dentures took place after Charles Goodyear invented the material in 1839. In 1843, Giovanni d’Arcoli showed how to use gold foils in filling cavities.

These are some of the very important milestones of dental materials developments, especially in restorative dentistry, and laid the foundation for the further developments of science and technology, as we know today.

14.3 Metals for Dental Application

Metals are used in dentistry because of their strength and durability. However, metals lack proper appearance and chemical durability. The current status is that virtually all metallic materials are suitable for use with minimal risk. However, complete elimination of “toxic” elements from all dental materials should be an everlasting goal.

14.3.1 Amalgams

Since 1895, dental amalgams have been used by dentists as direct filling materials [5]. Thus, these materials have a proven track record of more than 200 years. In spite of concerns about the presence of Hg, it is still safe to use, easy to manipulate, inexpensive, and durable. This is in spite of the competition from composites and glass-ionomer restorative materials. Some patients prefer amalgams to other alternatives because of their safety and cost-effectiveness.

Dental amalgams are formed by chemical reactions of Hg with intermetallic compounds in the Ag–Sn system as well as Cu. The reaction of liquid Hg with solid particles is known as amalgamation (also known as trituration). Mercury makes up about 45–50% of a typical amalgam composition. Once the amalgamation reactions are complete, the mixture is placed into the patient’s teeth under pressure. The mixture solidifies and within a few minutes the amalgam increases in strength and hardness. In an ideal situation, an amalgam should retain its dimensions throughout the restoration period. However, in reality this is not the case. A variety of factors lead to dimensional changes causing several complications. For example, contraction results in microleakage, plaque accumulation, etc.

The various important phases that take part in the amalgamation process are shown in Table 14.1. In a typical low copper composition, the following reaction occurs [6]:

$$ \gamma +\mathrm{Hg}\to {\gamma}_1+{\gamma}_2+\gamma \left(\mathrm{remnant}\right) $$

Table 14.1 Symbols and stoichiometry of phases that are important in dental amalgams [5]

The microstructure consists of γ embedded in the matrix of γ₁ and γ₂. Cu and Zn, although present in minor amounts, play important roles. Cu improves the physical properties of amalgams and hence high Cu-containing compositions were developed [7]. The high-Cu containing amalgams do not contain γ₂ phase. This is because Sn preferentially reacts with Cu to produce the η phase.

The shape of the original γ particles and Cu content classify the amalgams. Depending on how they are processed, particles are referred to as lathe-cut (elongated particles) and spherical (produced by the atomization process). These are available in traditional as well as high copper compositions. Finally, the lathe-cut and spherical powders can be admixed. The properties of such compositions are given in Table 14.2. Most of the modern alloys are high Cu containing alloys.

Table 14.2 Physical properties of dental amalgams [5]

The reaction between Hg and γ particles is performed in mechanical amalgamators. These motorized machines vibrate at a very high speed to mix the Hg and γ particles placed within a vial. For proper mixing, it is important to pay attention to the amalgamation time so that all the free Hg is consumed. Once the proper mix is obtained, it is condensed into the properly prepared cavity. Either hand condensation or mechanical condensation can be used.

As an alternative to the Hg-containing amalgams, gallium alloys were developed using the fact that Ga is a very low melting point metal [8]. Ga can be amalgamated with silver-tin-copper alloys in the same conventional fashion. The commercial compositions contain significant amounts of Pd to improve the corrosion properties. After triturating, the resulting microstructure consists of CuGa₂ and PdGa₅ surrounding the unreacted alloy particles within a mixture of Ag₉Iu₄.

14.3.2 Biocompatibility of Dental Amalgams

The usage of dental amalgam has remained a controversial issue. According to the U.S. Public Health Service, dental amalgams do not possess any immediate health hazard. However, studies found that mercury amalgams are unstable due to mercury’s low vapor pressure and galvanic action, resulting in leaking mercury vapor continuously into the lungs and saliva [9]. There is no evidence in the published literature that shows that the minute amount of mercury can cause poisoning to patients, except for those who have allergic reaction to mercury [10].

14.3.3 Casting Alloys

Along with amalgams, various casting alloys play an important role in restorative dentistry. From a historical viewpoint, Taggart introduced the cast inlay fabrication in 1907 [11]. He practiced the “Lost Wax (Investment)” casting technique, and since then, it remains an important technique for the fabrication of restorations. The casting alloys can be subdivided into all-metal restorations, porcelain fused-to-metal (PFM) restorations, and soldering alloys.

The alloys can also be grouped in terms of their compositions. American Dental Association (ADA) classifies them as high noble (HN), noble (N), and predominantly base metal (PB) [12]. According to the periodic table, the eight noble metals are gold, silver, and the platinum group metals. However, in an oral environment, silver is not necessarily considered noble. According to the ADA specifications, HN must contain more than 40 wt% of gold and the rest can be either gold or platinum group metals. Noble group should contain more than 25 wt% of noble metals, while predominately base metals (PB) should contain <25 wt% of noble metals. Table 14.3 describes various alloy systems in all-metal and PFM restorations. This discussion will begin with NB and N and ending with PB materials.

Table 14.3 Classification for full-metal and PFM restorations

For use in restorative dentistry, the casting alloys must satisfy many different criteria. These criteria include biocompatibility, corrosion resistance, tarnish resistance, allergenic reactions, esthetics, thermal properties, melting temperatures, extent of shrinkage, strength requirements, castability, bonding to porcelain, and economical aspects.

The casting alloys are further classified according to their mechanical properties. According to the ISO/DIS 1562 standard for casting gold alloys, several types of alloys are proposed [12]:

• Type I (low strength) – for making castings subjected to very low stress (in inlays), with minimum yield strength of 80 MPa and a minimum p.c. elongation of 18%

• Type II (moderate strength) – for castings subjected to moderate stress (inlays, onlays, and full crowns) with a minimum yield strength of 180 MPa, and minimum p.c. elongation of 10%.

• Type III (high strength) – for castings subjected to high strength (onlays, thin copings, and full crowns), the minimum yield strength is 270 MPa, and the minimum p.c. elongation is 5%.

• Type IV (extra high strength) – for castings subjected to very high stress (partial denture framework) with minimum yield strength of 360 MPa and the minimum p.c. elongation of 3%.

Types I and II are mostly used for inlays and their use is becoming lesser with the esthetically appealing natural looking materials. Types III and IV have the capability to strengthen due to several solid state transformations resulting in their high strength.

As shown in Table 14.3, all-metal restorations do not have the characteristics necessary for the PFM applications. This is because PFM-capable alloys must satisfy additional criteria in order to be compatible with porcelain. These require that: (1) the alloys melt at temperatures above 1100 °C; (2) they form thin stable oxides to promote atomic level bonding with porcelains; (3) they are compatible with thermal expansion coefficients in metal and porcelain (between 12.7–14.8 × 10⁻⁶/°C for alloys and 10.8–14.6 × 10⁻⁶/°C for porcelains; (4) they minimally creep or sag during firing of porcelains.

As seen in Table 14.3, at least three different classes of predominantly base metal (PB) alloys are used. These are cobalt-chromium, nickel-chromium, and titanium alloys. The former two are mostly used in metal-ceramic restorations and rarely used for all-metal restorations. All of these alloys were developed for restorative dentistry applications in the mid-1970s when the price of gold increased rapidly. Chromium provides tarnish and corrosion resistance of these alloys. Cobalt and nickel increases the Young’s moduli of Ni-Cr alloys and can be further sub-divided into beryllium-containing alloys and those containing no beryllium. Be addition decreases the melting point of alloys, which is useful for PFM applications.

14.3.3.1 Titanium and Related Alloys

Titanium has widespread uses in dentistry, which necessitates a separate section. Titanium is light, strong, ductile, and corrosion resistant [13]. Titanium’s high strength-to-weight ratio has been a major reason for titanium’s popularity. In the early 1950s, the newborn jet aircraft industry led to the rapid growth in titanium development technology. Starting in the 1960s, industrial uses of titanium began to become significant. Titanium was found to be one of the most “corrosion-resistant” structural metals, and marine applications such as power plant heat exchangers and water jet propulsion systems started using them [13].

Health-conscious patients and dentists alike are becoming more and more demanding in their quest to attain the “smile of the future.” Both are demanding a high degree of comfort and 100% organic compatibility. While titanium can match these demands, other dental alloys can cause unpleasant side effects of chemical and physical reactions in the mouth. Thus, modern dentistry is increasingly turning to titanium as an alternative to other conventional dental alloys.

In comparison to gold alloys, titanium yields a wider range of strength properties at a specific gravity that is four times less than gold. These physical properties allow for stronger and lighter crown and bridge frameworks. For the sensitive mouth, titanium is the best solution, and is especially suitable for patients suffering from allergies. Because of its low weight, the new titanium prosthesis causes little inconvenience to the patient. Its low thermal conductivity means that hot and cold foods can be enjoyed without discomfort. For the patient, these benefits add up to a new confidence level, providing a greater quality and enjoyable lifestyle.

Titanium has been established to be “biocompatible” with the human body and is classified as “hypoallergenic .” Titanium is totally resistant to attack by bodily fluids and is not recognized as a foreign material when utilized in implants such as artificial hips, pins for setting bones, and other biological implants. In dentistry, titanium actually encourages bone to grow to it, and the rapid growth in popularity of dental implants can be attributed in large part to titanium. With the application of unique, new dental porcelains to titanium, dentists can prescribe the most advanced and biologically safe porcelain to metal crowns and bridges [12].

Commercially pure (CP) with 99% titanium, and small quantities of Fe, C, O, N, H has five different grades (1–4 and 7) depending on the oxygen and iron content. The strength goes up from grade 1 to grade 4 [12, 13]. The most significant advantage of Ti is the presence of a thin passivating oxide film on its surface, which provides it with a very high corrosion resistance. However, the most important disadvantages of Ti are its reactive nature at the high melting point of 1668 °C, which needs special atmosphere during casting.

Titanium undergoes phase transformations from α phase (HCP) at 885 °C to β phase (BCC) structure. Alloying additions stabilize various phases. While vanadium stabilizes the β phase, aluminum stabilizes the α phase. Thus the most important alloy Ti–6Al–4 V is an α–β alloy. Table 14.4 lists the physical and mechanical properties of Ti and representative alloys.

Table 14.4 Physical and mechanical properties of Ti and its alloys

As noted before, an oxide layer (TiO₂) of 3–5 nm that forms on the surface makes Ti and its alloys extremely corrosion resistant. This is called the passive condition of titanium. Both CP titanium and its alloys have been shown to be biocompatible. The chemical inertness of implant material plays a very important role in long-term clinical success of osseointegrated implants. It is well known that most implant materials leak ions into the tissues. Some undesirable biological effect include: (1) alteration of fibroblast and calcium deposition during osteogenesis; and (2) peri-implant bone formations and the deposition of calcium ions on titanium surface.

14.3.3.2 Casting and Soldering

Other than the soldering alloys , the casting alloys are investment cast. The investment casting process can produce the fine features of restorations with high-dimensional accuracy. ANSI/ADA specification 2 describes several types of investments (molds) into which alloys can be cast. The investments can be made of gypsum or phosphate bonded. Different types of casting machines are available for filling the mold. In laboratories, either a torch flame, or induction/resistance heating can be used to melt the alloys, followed by centrifugal casting of the molten alloys into a cavity. Alternatively, vacuum melting and Ar pressure casting can be performed. After casting and solidifying, the investment is removed, and the casting is given the final shape. For achieving desired mechanical properties, the castings can be further processed . Types I and II alloys do not harden, while Types III and IV alloys are hardenable alloys. Typical heat-treatment of alloys requires heating to temperatures between 200 and 450 °C followed by quenching in water.

Soldering enables a technician to join simple cast parts. Dental casting alloys that can be soldered are both noble and base metal alloys. The term “presoldering” is used for metals prior to firing the veneers and “postsoldering” is used after veneering has taken place. Like the conventional soldering process in structural materials, soldering in dentistry needs filler and a flux, besides some suitable heat source. They are supposed to perform the exact tasks that their structural counterparts do. Gold-base fillers are normally used for noble metal alloys and silver-base fillers are used for base metal alloys. Traditional heat sources such as a torch can be used for soldering. Various lasers are now available for soldering Ti-alloys.

14.3.4 Wrought Alloys as Orthodontic Wire

Orthodontic wires are used in order to correct misaligned teeth. Usually, wires are made up of wrought alloys. As opposed to castings, they are usually cold worked and eventually formed by mechanical processes such as rolling, extrusion, and drawing.

One of the requirements of orthodontic wire is that they must have low elastic modulus, high yield strength, and good corrosion resistance. Besides, materials have to be sufficiently ductile for easier fabrication. Specifically, most wires should easily be soldered or brazed together. Four major alloys of orthodontic wires include austenitic stainless steels, cobalt–chromium–nickel, nickel–titanium, and beta titanium. Among these, even though stainless steel and cobalt chromium nickel alloy wires are far less expensive than Ti–Ni and beta Ti wires, former materials are not so popular. Table 14.5 describes the various materials and their properties.

Table 14.5 Wrought alloys and their properties

The shape memory TiNi alloys call for some additional discussion because of their interesting and beneficial properties for orthodontic applications. These alloys are associated with such properties as “Shape Memory” and “Superelasticity.” Buehler and Wang discussed the shape memory effect in an equimolar composition of Ti and Ni [14]. The shape memory effect refers to an alloy’s ability to return to a previously defined shape when heated according to appropriate schedules . While many compounds with shape memory properties have been identified, TiNi is in the forefront in orthodontic applications [15]. The diffusionless martensitic transformations are able to generate high strains and/or forces in TiNi. In addition, TiNi is biocompatible.

Four temperatures usually describe the SMA effect in TiNi: martensite start (M_s), martensite finish (M_f), austenite start (A_s), and austenite finish (A_f). The low-temperature martensitic phase is soft and deformable. As the temperature is increased, there is diffusionless transformation to the high-temperature austenitic phase. The thermal cycling between these phases actually creates hysteresis. Furthermore, SMAs can result in superelasticity with very high recoverable strains as compared with traditional elastic materials. It is critical to pay some attention to the superelastic behavior. This requires deformation to occur by martensitic transformation rather than by plastic deformation. In TiNi, plastic deformation is more dominant between 100 and 150 °C. Based on thermal and mechanical measurements, a superelastic window in temperature can be obtained, where there is minimal remnant permanent strain, resulting in ideal superelasticity. The characterization of SMAs requires precision determination of aforementioned transformation temperatures using differential scanning calorimeter (DSC).

In the past, the shape memory alloys were given shape at high-temperature exposure. After cooling down, the clinician would manipulate the wire in the patient’s mouth and it would get into the original shape of the wire. The superelastic wires have much more beneficial springback (the ratio of yield strength to elastic modulus) than their competitors. More recent alloys contain some amount of copper. The addition of Cu approximately 5 wt% instead of Ni and 0.29 wt% of Cr makes the transformation purely shape memory without the presence of the intermediate-R phase and shows superelastic behavior at 37 °C, the temperature of the human body.

As discussed before, β-titanium orthodontic wires have been used as orthodontic wire. β phase is stabilized by the addition of Mo [16]. The favorable springback behavior and formability are also advantageous.

Although Ti–Ni shape memory alloys and 316 L austenitic stainless steel are used in many dental applications, they contain Ni, which causes metallic allergy. Non-allergenic nickel-free stainless steel, in which mechanical strength and corrosion resistance are improved using nitrogen as a substitute for nickel, has drawn attention as a new biomaterial. However, because the material itself becomes extremely hard, depending on the forming and machining conditions, Ni-free stainless steel dental devices and technologies for manufacturing small-scale, complex-shaped dental devices from Ni-free stainless steel have not reached practical application. Technology development in this field is strongly desired.

14.3.5 Dental Implants

A dental implant is an artificial tooth root (synthetic material) that is surgically anchored into the jaw to hold a replacement tooth or bridge in place [17, 18]. The benefit of using implants is that they do not rely on neighboring teeth for support. They are permanent and stable. Implants are a good solution to tooth loss because they look and feel like natural tooth. Implant material is made from different types of metallic and bone-like ceramic materials that are compatible with body tissue. Today, dental implants are becoming more predictable. Studies have demonstrated success rates ranging from 80% to 92% for the maxilla over 5–10 years. Advances in surgical procedures, such as maxillary sinus lifts, lateral inferior alveolar nerve placement, and guided bone regeneration (GBR) allow the practitioner to place implants in patients that would not have received this treatment option 20 years ago [19].

In general oral implants can be categorized into three main groups:

• Endosseous implants

• Subperiosteal implants

• Transosseous implants

14.3.5.1 Endosseous Implants

Endosseous implants are the most frequently used implants today [20]. Based on their shape, function, surgical placement, and surface treatment, they could be further categorized into several sub-categories. However, for the present purpose, only three main categories will be described. The three main categories are root-form, ramusframe, and blade-form.

Endosseous root-form implants are surgically inserted into the maxilla or mandible, whereby the implant body generally penetrates the cortical bony plate. In 1939, Strock first attempted to change the shape of a dental implant from that of a rootlike tooth to an implant with a threaded body resembling a wood screw [3]. Recognizing the need for implants to be biologically compatible, he made the implant from an alloy of chromium–cobalt–molybdenum (called Vitallium). He found that certain metals, when in contact with tissue fluids, produce a galvanic action. This ultimately corrodes the metal appliance. Vitallium was the only metal used at that time that produced no electrolytic activity when implanted into biological tissues. Among these, Strock’s implants are the earliest documented long-term endosseous dental implants. Furthermore, he appreciated the need to direct axial forces and minimize nonaxial forces to the dental implants. Root-form implants are considered by some as the standard of care in oral implantology [21]. The implant, a screw-shaped form like the original root, is screwed into the patient’s jawbone. The screw’s shape and size are based upon the tooth being replaced and/or the size and shape of the remaining jawbone. The implant is generally made out of titanium alloys, aluminum oxide, Vitallium, commercially pure titanium or sapphire [21]. Once the implant has been placed into the jawbone, a few months are required for the implant and the bone to fuse. This fusing creates a solid, secure implant to which an abutment can be attached. Before the prosthesis phase, a temporary crown or healing abutment is attached to the implant for up to 8 weeks. Gum tissue grows around the temporary crown. After the implant has fused, an impression of the tissue shape and the implant is taken. From this impression a permanent tooth is fabricated. The permanent-fabricated tooth is then cemented to a custom abutment that is then screwed onto the implant. Root-form implants are the closest shape and size to the natural tooth root. They are commonly used in wide, deep bone to provide a base for the replacement of one, several, or a complete arch of teeth.

Ramusframe implants belong in the category of endosseous implants. These implants are designed for the toothless lower jaw only and are surgically inserted into the jaw bone in three different areas: the left and right back area of the jaw (the approximate area of the wisdom teeth), and the chin area in the front of the mouth.

In 1968, Linkow introduced a flat titanium endosteal blade implant , which often serves as a means of using the narrow and/or shallow areas of remaining alveolar bone where dimensions do not permit the use of the cylindrical or root-shaped implants [3]. These implants primarily form a connective-tissue interface, a process called fibro-osseus integration. Connective tissue forming around the implant takes the form of collagenous fibers that are parallel to the dental implant surface. However, since these are also surgically placed into the bone, we categorize them into the endosseous implant category. Blade implants have a long track record, much longer then the root-form implants. Their name is derived from their flat, blade-like (or plate-like) portion, which is the part that gets embedded into the bone. Blade implants are not frequently used anymore; however, they do find an application in areas where the residual bone ridge of the jaw is either too thin (due to resorption) to place conventional root-form implants.

14.3.5.2 Subperiosteal Implants

The second type of dental implant is the subperiosteal , or on-the-bone, which has been used successfully for the past 30 or 40 years [22]. It incorporates a rigid, plate-like element and is often used for areas without adequate bone for cylindrical endosteal implants. Subperiosteal implants are those that typically lie on top of the jawbone, but underneath the gum tissues. The important distinction is that they usually do not penetrate into the jawbone.

Subperiosteal implants were introduced in the 1940s [3]. Of all currently used devices, it is the type of implant that has had the longest period of clinical application [22]. Unlike endosseous implants, these implants are not anchored inside the bone. Instead they are shaped to ride on the residual bony ridge of either the upper or lower jaw. Subperiosteal implants have been used in completely edentulous (toothless) as well as partially edentulous upper and lower jaws. However, the best results have been achieved in the treatment of the edentulous lower jaw.

While infection is a potential risk with subperiosteal implants , loosening over time is another problem and must be replaced in the upper jaw after 5 years. In the lower jaw, these implants last up to 10 years. Replacing these implants also requires a surgical procedure.

14.3.5.3 Transosseous Implants

Transosseous implants similar in definition to endosseous implants in that they are surgically inserted into the jawbone [22]. However, these implants actually penetrate the entire jaw so that they actually emerge opposite the entry site, usually at the bottom of the chin. This is also the site where they are secured with a device similar to a nut and a pressure plate.

14.3.5.4 The Phenomenon of Osseointegration

In 1952, Dr. Per-Ingvar Brånemark, an innovative orthopedic surgeon, embarked on a series of basic pioneering studies unrelated to implants that could define the conditions necessary for regeneration of highly differentiated tissue after injury [23]. Using cylinders made of titanium to examine the repair mechanisms of bone, Brånemark reported that a direct functional and structural bond had occurred between organized vital bone and the inanimate alloplastic material. He called this process “Osseointegration.” He suggested that osseointegration occurred only with the titanium and not with cylinders made of other materials. On the basis of these results, he designed and executed a series of investigations using titanium as a cylindrical threaded endosteal implant for replacing a single tooth or anchoring fixed bridgework [24].

By 1965, Brånemark’s implants were first placed in patients. They consisted of a cylindrical screw-type implant made of pure titanium and induced osseointegration between titanium metal and the peri-implant bone. Degradation of the titanium in vivo is minimal because of a protective oxide layer that develops between the titanium implant and the bone. With its provisional acceptance by the American Dental Association in 1986, this innovation has greatly contributed to the acceptance, resulting in widespread use of dental implant treatment by the dental community [25].

Osseointegration is a mode of implant attachment with a bone anchorage, in contrast to a soft tissue-attachment implant. The former interface is rigid, whereas the latter tends to permit macroscopic movements. Various histology-based definitions of osseointegration have not reached a general consensus because of difficulties in properly defining the resolution level of the direct bone-to-implant contact and the proportion of a bony contact necessary to call the implant osseointegrated [26, 27] (Figs. 14.1, 14.2, and 14.3).

Fig. 14.1

(a–c) Various forms of endosseous implants. (a) Root-form. (b) Ramus-form. (c) Plate-form [23]

Fig. 14.2

Subperiosteal implant [23]

Fig. 14.3

Transosteal implant [23]

Osseointegrated implants are most stable because they allow bone tissue to grow around them. Osseointegrated implants are made of titanium. Studies indicate that the body’s tissues may better tolerate titanium than the chrome-cobalt metal that comprises other implants [28]. Because the bone thickens and grows attached to these implants, they become stronger with time and cause little or no bone re-absorption. Unlike other implants, they feel and function like natural teeth [29, 30].

As better scientific analyses and more clinical experience have been acquired, the definition of osseointegration evolved. A convenient working clinical definition is more useful and the term is currently best defined as a process whereby clinically asymptomatic rigid fixation of alloplastic materials is achieved, and maintained in bone during functional loading [31]. Osseointegration was the hallmark of success in implant dentistry in the 1980s. It was believed that an implant was successfully integrated when there was direct contact between bone and the titanium implant (at the light microscopic level) with no fibrous connective tissue interface.

14.3.5.5 Materials Issues in Dental Implants

Ti, Ti–Al–V alloys, Co–Cr–Mo alloys and stainless steel are easy to manufacture and shape, have adequate physical properties, and can integrate with bone. That is why metals are the most common implant materials [13]. The major disadvantages are corrosion, fatigue failure, and adverse stress distributions. The material of choice is titanium or Ti–V–Al alloy. Common misconceptions are that implants are made of “pure” Ti and they do not corrode. It is extremely difficult, and economically unviable, to remove oxygen from Ti. Thus the most common metal is “commercially pure” Ti. As noted before, the passive layer is typically 15 nm thick and prevents the metal from further corrosion. The passive layer on Ti and its alloys is unique in that it is stable under all normal physiological conditions. Furthermore, this layer is involved in the integration of the implant with bone. The favored alloy is Ti-6% Al-4% V because it has a low density, an elastic modulus half that of cobalt-chrome or stainless steel, has adequate fatigue properties, and is biocompatible. Al and V are phase stabilizers, and the alloy can be heat treated to modify its properties. The major disadvantage is that it has a low torsional strength that can result in fracture of screw implants. These alloys are often said to osseointegrate with bone.

14.3.5.6 Surface Issues

Surface features that need consideration to improve material-tissue interaction include material purity, surface chemistry, surface charge, surface morphology, and surface coatings. Data demonstrate that rough implant surfaces have increased bone-to-implant contact. Greater forces to break the bone-implant interface compared to more smooth surfaces. In clinical trials, rough surfaced implants had significantly higher success rates compared to implants with more smooth surfaces. Various surface treatments that will be considered in the following section include etched titanium, titanium plasma spray (TPS) , hydroxyapatite-coated (HA-coated) titanium, bioglass coating, and biomimetic surface modifications [44].

Etched Surface – Titanium surface is etched with acid. This surface demonstrates increased roughness and therefore greater surface area which is associated with increased dispersion of blood cells around the implant body. An increase in blood-cell dispersion around an implant body has been shown to facilitate osseointegration. SEM and laser analyses have demonstrated the effectiveness of the acid-etching procedure by showing the superior roughness and therefore increased surface area achieved [32, 33].

Titanium Plasma Spray (TPS) – Titanium plasma spray (TPS) coating on implants is a pure titanium powder that is applied to the implant using an advanced vacuum plasma-spray process. TPS-coated implants have a relatively large surface area. TPS-coated implants have been shown to possess exceptionally high mechanical retention properties in bone (Fig. 14.4).

Fig. 14.4

(a–c) Various surfaces of titanium used in dental applications. (a) Acid-etched surface. (b) TPS-coated surface. (c) HA-coated surface [37]

The most common technique to produce porous surface is plasma-spraying titanium powder on the implant to create macroporous titanium surface layer [34]. The plasma-sprayed surfaces were reported to have advantageous morphological features that make it more suitable to successful clinical applications. However, this technique has several disadvantages. First, the geometry of the interface between the porous coating and the substrate were found to cause stress concentrations that reduced the fatigue strength [35]. Second, high-temperature heat treatments often used to bond the porous coating and substrate were attributed to changes in microstructure and eventual reduction in strength [36]. Lamellar structure, micropores, and microcracks have also been observed inside the plasma-sprayed titanium coating [37].

Hydroxyapatite-coated (HA-coated) titanium  – Studies demonstrated that the osseoconductive properties of HA-coated implants significantly accelerate osseointegration. This ensures a good tissue response and an optimal chemical bond between the coating and the bone. Furthermore, the increased roughness of the HA coating created by a plasma spray process helps lock the implant into the bone.

Excellent short-term clinical results were reported for hydroxyapatite-coated implants [38, 39]. HA-coated implants have higher integration rate, promote faster bone attachment, and achieve direct bone bonding with higher interfacial attachment strength to none when compared to non-coated implant [40]. However, there are many controversies regarding their long-term prognosis. An 8-year-old clinical study found that the survival rate was initially higher for HA-coated implants, but decreased significantly below that of titanium plasma-sprayed implants after 4 years [40]. Most of the delayed failures were associated with inflammatory disease. HA delaminating was obtained in both animal experiments and human clinical experience [41].

The resorption of the HA coating compromises long-term implant survival. The coating often contain impurities such as calcium phosphates and CaO, which decreases chemical stability and enhances degradation of the coatings. Delamination of the coatings is due to fatigue and/or thermal expansion coefficient mismatch between Ti/HA interface.

Bioactive Glass Coating – Bioactive glass is a bioactive material that promotes the formation of bone in the human physiological environment. Some bioactive glasses also form bonds with soft collagen-based tissues. One barrier of using bioactive glass as coating is the high thermal expansion coefficients (14–15 × 10⁻⁶/^oC). Tomsia and his colleagues demonstrated that bioactive glass could be tuned to the same thermal expansion coefficient as the Ti alloys it coats [42, 43]. Glasses in the system Si–Ca–Mg–Na–K–P–O were successfully applied on Ti6Al4V using conventional enameling technique. The thermal expansion coefficients obtained were close to that of Ti6Al4V of 9.6 × 10⁻⁶/°C. The bonding between the glass and coating was achieved by the formation of a Ti₅Si₃ interlayer. Further discussion of bioglass can be found in Sect. 14.4.6.

Biomimetic surface modifications  – Discoveries by Kokubo and his colleagues indicated that chemically treated Ti and its alloys are bioactive. After being subjected to 5.0 M-NaOH treatment at 60 °C for 24 h and subsequent heat treatment at 600 °C for 1 h, they are soaked in simulated body fluid (SBF) [44, 45]. Titanium-based alloys such as Ti6Al4V, Ti6Al2Nb, and Ti15Mo5Zr3Al formed the bone-like apatite layer on the surface in SBF. This bonding accelerates the stabilization of implants and extends its lifetime before failure. The apatite formation on the surfaces of Ti and its alloys was assumed to be induced by a hydrated titania which was formed by ion exchange of the alkali ion in the alkali titanate layer and the hydronium ion in SBF. Through in vitro studies Kokubo and colleagues showed that the alkali-treated Ti provides the most favorable conditions for differentiation of bone marrow cells [46] (Fig. 14.5).

Fig. 14.5

Titanium surface and the formation of apatite as suggested by Kokubo [44]

In vivo studies conducted in rabbit tibiae to determine the effects of chemical treatments and/or surface-induced bone-like apatite on the bone-bonding ability of titanium implants. Biomechanical examination results showed that treated implants exhibited significantly higher failure loads compared with untreated Ti implants at all time periods. They revealed that treated Ti implants directly bonded to bone tissue during the early post-implantation period, whereas untreated Ti implants formed direct contact with the bone only at 16 weeks [47,48,49,50]. The results of their study suggest that chemical treatments accelerated the bone-bonding behavior of Ti implants and enhance the strength of bone-implant bonding by inducing a bioactive surface layer on Ti implants. They concluded that alkali- and heat-treated titanium offers strong bone bonding and a high affinity to bone as opposed to a conventional mechanical interlocking mechanism. Alkali and heat treatments of titanium may be suitable surface treatments for cementless joint replacement implants. These results imply that the Ti–OH groups on titania, proposed to be responsible for the apatite formation, are effective for apatite nucleation when they are arranged in a specific structural unit based on the anatase structure. Alkali- and heat-treated implants also showed direct bonding to bony tissue without intervening fibrous tissue [50]. On the other hand, untreated implants usually had intervening fibrous tissue at the interface between bone and the implant. The early and strong bonding to bone of alkali- and heat-treated titanium and its alloys without intervening fibrous tissue may be useful in establishing cementless stable fixation of orthopedic implants.

14.3.5.7 Problems with Dental Implants

Clinically, lack of osseointegration leads to implant mobility and subsequent failure. Therefore, a mobile implant is a failed implant. Non-functional, failed implants must be removed to prevent the associated bone loss from continuing. For long-term successful performance of all dental implant types, the following general factors should be considered. These are biomaterials, biomechanics, dental evaluation, medical evaluation, surgical requirements, and healing processes.

Implants fail for a variety of reasons. Some studies have related failures to biological or microbiological factors, while others attribute dental implant failure to biomechanical factors, biomaterial factors, or implant surface treatments and characteristics [32, 33]. Improper patient selection is a significant reason for failure. Obviously, a patient unmotivated to control plaque around natural teeth would not be a good candidate for dental implants. Implants are doomed to fail when placed in patients having insufficient quality and/or quantity of bone to support the implant fixture.

Systemic health of the patient is important when considering implants. Recent studies have demonstrated similar failure rates between well-controlled diabetics and non-diabetic controls or only slightly higher failure rates with Type 2 (non-insulin dependent) diabetics [34]. Uncontrolled diabetic patients are poor candidates for any surgical procedure. Inferior surgical technique is another possible cause of implant failure. During surgical placement, implant failure may result from inadequate irrigation of the surgical site or from using low torque and excessive drill speed during placement. Failure results from excessive temperature elevation in bone during placement, leading to necrosis of the supporting bone around the implant. Inadequate implant restorations may also contribute to implant failure. Restorations placed on endosseous implants may cause traumatic occlusion, leading to failure. Additionally, poorly restored implants may have overhangs or be over-contoured, which may lead to plaque accumulation and eventual failure.

The main factors in implant failure are infection and occlusal force stresses. The microbiota around stable vs failing implants seem to parallel the patterns observed around healthy vs diseased natural periodontal sites. Recent studies indicate that microorganisms associated with periodontal disease are found in higher proportions in failing implant sites. Periodic plaque removal and health maintenance should help prevent such failures.

With single-implant therapy, survival also depends on the position of the implant in the mouth. Some reported on a 24-month life-table analysis study on two-stage implant survival, using a variety of implant designs [32, 33]. They reported survival to be 100% in the anterior mandible, 92% in the posterior mandible, 94% in the anterior maxilla, and 78% in the posterior maxilla. The lower survival rate in the latter area is probably the result of the cancellous nature of the bone and thin cortical plates. Small-diameter cylinders appear to do less well than large diameters or blades in hollow, cancellous bone, and thin cortical plates. Various implant shapes are needed to overcome bone morphological limitations and meet restorative requirements. Unfortunately, implants for partially edentulous patients are most often needed in the posterior mandible and maxilla, reflecting the pattern of natural tooth loss.

14.4 Ceramics for Dental Applications

As opposed to metals, the strength and toughness values of ceramics are not as good. In spite of that, there is a long history of using ceramics in restorations. The two main reasons for the use of ceramics in dentistry are the ability of restorations to match the esthetics and functions of missing teeth [51,52,53,54]. Furthermore, ceramics are benign and do not result in any cytotoxicity.

It is obvious that ceramics can match the opalescent appearance of teeth that no metal can provide. Ceramics can also withstand the cyclic masticating forces that vary between 400 and 800 N. They can also provide the right balance and support between bone/soft tissue. While it is known that ceramics are less mechanically reliable than metals, more and more interesting developments are taking place in ceramics, resulting in higher mechanical strength and toughness values [55, 56].

Most common uses of ceramics are inlays and veneers. In these simple applications, strength and fracture toughness values are less important than their esthetic attributes. Ceramics are also used as crowns, a structural replacement of a tooth. These are thin-walled structures, which may be composed of metal-ceramic or all-ceramic materials. Some examples are depicted in Fig. 14.6. Like the various applications of metals in dentistry, dental ceramics can be classified according to several schemes, which may include their actual use, composition, processing methods, firing temperatures, microstructure, mechanical, and esthetic properties. Instead of classifying them in various sub-sections, it is simpler to categorize them into metal-ceramic and all-ceramic systems. The metal-ceramic systems are essentially feldspathic porcelains, while all-ceramic systems consist of glass-ceramics, glass-infiltrated ceramics, and fully sintered monolithic ceramic materials. This simple classification can be further subdivided according to processing routes, indicated uses, microstructure, esthetic signatures, and mechanical properties. It is also worth mentioning that dental ceramic literature often refers to various commercial products. Therefore, as and when appropriate, these commercial names will be referred to in the text. (Table 14.6).

Fig. 14.6

(a–d) Types of dental restorations. (a) In-Ceram crown with standard crown design. (b) Lateral view of In-Ceram crown demonstrating underlying tooth structures. (c) Lateral view of ceramic crown on a titanium implant. (d) Porcelain veneer resin bonded to tooth enamel

Table 14.6 Strength and toughness of tooth structures and dental ceramics [57]

14.4.1 Metal-Ceramic Restorations

The feldspathic porcelains , required for such uses, are based on patents by Weinstein et al. [58] and Weinstein and Weinstein [59]. The patents are about developing compositions with tailored thermal expansion coefficients and the process of their bonding to metallic alloys. The compositions are based in the SiO₂–Al₂O₃–K₂O ternary system. Specifically, SiO₂ and potash feldspar (K₂O·Al₂O₃·6SiO₂) are the basis for dental porcelains. Other additives such as opacifiers (TiO₂, ZrO₂, and SnO₂), appropriate pigments, and fluxes (to lower melting point) are also added. The porcelains are normally classified in terms of their firing temperatures: high firing (1300 °C), medium firing (1100–1300 °C), low firing (850–1100 °C), and ultra-low firing (<850 °C).

The most important aspect of the development of feldspathic porcelain is the presence of leucite (KAlSi₂O₆). The presence of leucite increases the thermal expansion coefficient of porcelains to an extent that bonding between the porcelain and metal substrate can easily take place with matched thermal expansion coefficients. While it is easy to say that leucite can increase the thermal expansion coefficients of porcelains, it is not an equilibrium phase. Different thermal processing histories result in different cooling rates of leucite with concomitant variation in thermal expansion coefficients [60]. While slow cooling in a furnace increases the leucite content from 11 vol% to 56 vol% [61], an isothermal hold at 750 °C 4–16 min (simulating post-soldering treatment), the leucite content in commercial porcelain increases from 6 vol% to 21 vol% [62]. The various thermal histories can also lead to cracking of porcelains due to the thermal expansion coefficient mismatch [63]. Microcracking can also happen due to the cubic-to-tetragonal transformation of leucite upon cooling [64].

The flexural strengths of feldspathic porcelains are quite low (between 60 and 70 MPa). However, when bonded with the metal substructure, the flexural strength increases significantly to approximately 300 MPa [12].

14.4.2 All-Ceramic Restorations

Glass-ceramics – These materials are composites of glass and ceramics, which can be processed into complex shapes using glass-shaping technology. These materials are, therefore, also referred to as castable/pressable glass-ceramics. The ceramic particulates are formed in a glassy matrix by carefully nucleating and growing ceramic crystals with suitable heat treatments (also known as ceramming). The initial development of Dicor® glass ceramics took place in Corning Glass Works and was marketed by Dentsply International. Dicor® consists of a magnesium aluminosilicate glass with 45% fluoromica (K₂Mg₅Si₈O₂₀F₄) crystals. The presence of these crystals enhances the mechanical properties of these glass-ceramic materials. In spite of the advantages of enhanced processibility and the potential for enhanced mechanical properties, the extent of mechanical property enhancement was not substantial. Furthermore, Dicor® materials do not have the ability to attain internal color. Even though, esthetically, they are better than metal-ceramics, with the development of other glass ceramics, Dicor® materials are no longer available on the market.

Another important glass ceramic is Empress 1®, which is essentially a leucite-reinforced feldspathic porcelain. The controlled nucleation of leucite enhances the mechanical properties while being produced by the conventional investment casting technique. Empress 1® has flexural strength values about the same as Dicor®. The translucency of this material is slightly less than Dicor® but can be tolerated in restorations.

A major advancement in glass ceramic development is the Empress 2®, which is a lithium silicate–based system. With proper composition choice and heat treatment, the microstructure consists of 70% of elongated lithium-disilicate crystals [65]. This creates a highly filled microstructure. The manufacturer recommends these be used for inlays, onlays, veneers, etc.

Monolithic ceramics  – An example of this type of material is fully sintered alumina (Procera Allceram®). This material is 99.9% alumina and sintered at 1600 °C to produce a core, which is compatible with dental porcelain. While the mechanical properties are quite good, the sintering process results in a large shrinkage. Therefore, in order to obtain high-dimensional tolerance, CAD-CAM system is used. Once the sintered coping is obtained, a matched porcelain is placed on the coping. Procera Allceram® is limited to use as a single-unit restoration in the anterior or posterior regions.

Zirconia (ZrO₂)  – The most significant advancement in the applications of monolithic ceramics in dentistry is the introduction of zirconia (ZrO₂) ceramics in view of their superior mechanical properties than any other ceramic material [66, 67]. Their superior mechanical properties are achieved in zirconia because of unique martensitic (diffusionless) phase transformations that occur in the material and the resultant increase in toughness. As a result, zirconia-based materials are mostly reliable in different load-bearing applications.

The phenomenon of polymorphism of zirconia is well known. The oxide material exists in three different crystal structures at different temperatures – cubic (c), tetragonal (t), and monoclinic (m). Of these, the high-temperature phase is the cubic phase, which exists at temperatures beyond 2370 °C [68]. The room temperature phase is the monoclinic phase [68]. Zirconia exists in a metastable tetragonal phase at 1170 °C. It is the martensitic transformation from the tetragonal to the monoclinic phase that makes zirconia materials so interesting in load-bearing clinical applications. It is noted that the phase transformation from the t to the m phase is accompanied with a 5% increase in volume. If the transformation occurs uncontrolled, a zirconia component can fracture easily, an outcome that is undesirable. On the other hand, the metastable tetragonal phase can be made stable by adding suitable dopants to zirconia, based on different zirconia-based binary phase diagrams . These dopants are, therefore, also referred to as stabilizers. The addition of stabilizers enables martensitic transformations to take place in a controlled fashion during the propagation of a crack or defect, thus resulting in an increase in toughness. In other words, the materials become fracture resistant and reliable under load-bearing conditions. These materials are known to be toughened by the mechanism of “Phase Transformation Toughening” or simply “Transformation Toughening.” Important stabilizers that are well known to provide transformation toughening include calcia (CaO), yttria (Y₂O₃), magnesia (MgO), and ceria (CeO₂) [69]. These dopants (stabilizers) result in different types of toughened zirconia materials according to binary phase diagrams of ZrO₂ with CaO, MgO, Y₂O₃, and CeO₂. The former two dopants of CaO and MgO result in “partially stabilized zirconia (PSZ)”– a matrix of larger c-zirconia grains containing transformable t-zirconia grains . The latter two dopants give rise to “tetragonal zirconia polycrystals (TZP),”, wherein the entire micro- (or nano-) structure consists of transformable t-zirconia grains. Depending on the use of stabilizers, these are referred to as Y-TZP or Ce-TZP. From the processing aspect, PSZ-materials need high-temperature sintering and subsequent careful heat treatment to achieve desirable fracture toughness values. The TZP materials can be sintered at lower temperatures without the need for further heat treatment and hence industrially more sought after. Finally, yet another technically and clinically important material can be obtained by making composites of Al₂O₃ with ZrO₂, which has no miscibility with Al₂O₃. Al₂O₃ provides a matrix with high elastic modulus embedding transformable t-zirconia grains. These are commonly known as zirconia-toughened alumina (ZTA) [69].

This discussion justifies the popularity of TZP materials in the dental device industry. Comparing yttria and ceria doping, the former can produce the most esthetically pleasing restorations with mechanical reliability. More specifically, the addition of 3 mol % yttria in Y-TZP is reported to possess high bending strength above 1000 MPa and fracture toughness of about 6 MPa.m^1/2 [68,69,70]. Furthermore, researchers reported the development of translucent Y-TZPs with good mechanical properties and esthetic appearance for dental restorations [71]. Even though, CeO₂ doping can result in higher toughness values, its characteristic cream color led industries to move away from its use in dentistry. It is the combination of mechanical reliability and aesthetics along with the ease in machining by CAD/CAM (e.g., Lava) to produce complex geometries that these materials are so popular in dentistry. For instance, zirconia implant abutment has shown excellent survival rates after crown insertion for 5 years [72].

In spite of all of these advantages, the major limitation of Y-TZP materials is aging or low-temperature degradation (LTD) . Lawson has published a comprehensive review paper on the phenomenon, which induces t-m transformation at low-temperature exposures between 150 °C and 400 °C [73]. These transformations are accompanied with surface degradation including surface roughening, grain pull out, surface uplift, and microcracking [74]. This issue has attracted substantial attention when a number of Prozyr femoral head implants failed prematurely with a concomitant increase of the m-phase content [75]. In general, t-m phase transformation occurs in the presence of moisture or local stresses. In the oral cavity, LTD occurs due to the presence of saliva, heavy mastication loads, fluctuating temperatures, oral bacteria, pH changes, etc., which could be deleterious for the dental application [74].

Important parameters that affect the LTD behavior of Y-TZP is essentially related to the stability of individual tetragonal grains from transformation. This stability/instability behavior is related to the type of dopant and its content in conjunction with resultant grain size. According to thermodynamic calculations, each stabilizer has a critical grain size that triggers the t-to-m transformation. In the case of Y-TZP, this critical grain size is approximately 0.4 μm. If the final grain size is lesser than this critical size, there is a high chance that the grains will be more stable prior to the ageing treatment.

Masaki did not see the occurrence of LTD by ageing Y-TZP in hot air at 200 °C for 2000 h, in cases where the initial grain sizes were below 0.5 μm [76]. Chen and Lu found that the critical grain size (0.37 μm) is even lower in ageing in moist environment as opposed to aging in air (0.52 μm) at temperatures between 100 °C and 500 °C for various time periods [77]. This result is an indication of the effect of the presence of water molecules on the phenomenon of LTD . Thus, most studies on LTD focus on exposing Y-TZP samples under hydrothermal conditions, particularly in superheated steam conditions [73].

The data reported in the literature regarding the degradation of mechanical properties are somewhat confusing. This may be due to lack of reporting data in a comprehensive manner. Kim et al. aged Vita In-Ceram YZ at different temperatures (20 °C, 75 °C, 100 °C, 125 °C, 150 °C, 175 °C, 200 °C, and 225 °C) in distilled water for 10 h [78]. The results showed that flexural strengths decreased with the increase in the m-phase content at about 54% at 150 °C. At even higher temperatures of 175 °C, the monoclinic phase content reached saturation at 75%, while the flexural strength continues to fall. Borchers et al., on the other hand, reported that regardless of aging condition, there was an increase in the monoclinic content but without significantly affecting the flexural strength values [79]. Cattani-Lorente et al. reported that hydrothermal aging of Y-TZP (Lava) in steam at 140 °C for 7 days resulted in a 30% reduction in Young’s modulus and hardness [80].

There are reports of aging in different media as well. Kawai et al. aged commercial Y-TZP (Daiichi) manufactured by in three different solutions, i.e., distilled water, Hank’s solution, and lactic acid at 140 °C for 7 days [81]. An m-phase content of about 60% was measured in the Y-TZPs regardless of medium type. This result shows that solution compositions have negligible effect on the ageing behavior of Y-TZP. In an oral environment, constant exposure to heavy loads of mastication, fluctuating temperatures, pH changes, and oral bacteria could amplify the effect of aging compared to when used in orthopedic implants in the body [75, 84]. For example, Miragaya et al. observed a decrease in the flexural strength from 1322.9 MPa to 843.7 MPa for commercial dental Y-TZP (Lava Frame) after exposure in human oral cavity using intra-oral appliances for 60 days [82]. Similar degradation in terms of flexural strength occurred with the commercial dental Y-TZP (Lava Plus), where the strength decreased from 955.9 MPa to 603.2 MPa. The decrease of flexural strength is accompanied with the formation of monoclinic phase of Y-TZP (Lava Frame) (7.7%) and Lava Plus (4.7%). The authors explained that the decrease of flexural strength was due to the presence of ions and humidity in the human saliva.

Given the various parameters addressed in the above, it is important to discuss standardized evaluation of Y-TZP components for susceptibility to LTD. In the case of industrial ceramics, the aging test conditions were in accordance with ISO standard for medical zirconia and the Japanese industrial standard, which is at 180 °C in water, 1 MPa for 1 h. This condition has become a reference for the dentistry industry [74]. The recommended aging temperature for Y-TZP ceramics for dental and industrial applications is harsher because aging at higher temperature provides an aggressive environment to gauge the vulnerability of the zirconia to undergo the phase transformation. The relevant ISO standard 13,356, used to recommend 1 h of aging at 0.2 MPa pressure in an autoclave at 134 °C, which is equivalent to service condition of for 3–4 years in vivo at 37 °C. However, after many failures of so many Prozyr femoral heads, those recommendations seem to be inadequate [83]. It is now recommended that the accelerated aging in water be carried out at 134 °C, 0.2 MPa for 5 h, which resemble the aging condition of 15–20 years of implants (in water at 37 °C) in the orthopedic industry [74]. Furthermore, following the recommendation of Chevalier, the critical grain size was reduced from 0.6 μm to 0.4 μm. It is also important to prescribe the monoclinic content to be less than 5% [83]. Lucas et al. investigated the degradation of commercial Y-TZPs having different grain sizes according to the revised ISO 13356:2008 standard by exposing the samples in water vapor at 134 °C/2 bar for 5 h [84]. They found that after aging for 5 h, the samples with a grain size of 0.428 μm yielded 4.03% of m-phase, whereas sample with a larger grain size of 0.574 μm yielded a high m-phase of 6.64%.

In view of the state-of-the-art of Y-TZP research and its pros and cons, future research should be directed to explore alternative materials to replace Y-TZP. The first strategy is to use additional dopants such as of silica [85]. These co-dopants (along with yttria) may act against the LTD while preserving beneficial toughness and strength values. These solutions are being actively pursued currently within the framework of the Y-TZP system. The problem with yttrium is that being a trivalent ion, it creates oxygen vacancies that help hydroxyl group diffusion in the lattice, generating nucleation of the transformation by stress corrosion type mechanism. Oxygen from the moisture occupies the vacant sites and hydrogen occupies the adjacent interstitial sites. Some tensile stress is generated in surface grains due to the lattice contraction that destabilizes the tetragonal phase [86,87,88]. Therefore, anionic vacancies play a key role in the water diffusion rate and hence in the aging kinetics. When two yttrium ions (Y³⁺) replace two zirconium (Zr⁴⁺) ions in zirconium sub-lattice, they create one oxygen vacancy (V_o) in the oxygen sub-lattice.

A second strategy is to replace yttria with ceria as suggested by Chevalier [83]. As ceria is isovalent with zirconia, it does not create any vacancies upon substitution. It is surprising that there are very few reports on the development of Ce-TZP materials as dental ceramics in spite of exhibiting superior toughness (up to 20 MPa.m^1/2) and almost no aging in view of the argument is presented in the above [86,87,88]. Thus, there is still some interesting possibility to develop monolithic zirconia resistant to LTD. The backlash due to Prozyr failures and the new generation of zirconia-alumina composites promoted by major ceramic companies (i.e., Ceramtec or Metoxit) play against further development of monolithic zirconia-based products in orthopedics.

This fact leads to the third strategy by way of developing alumina-zirconia composites. This may be the way to benefit from zirconia transformation toughening without any LTD-related issues. In the recent literature concerning alumina–zirconia composites for biomedical applications, different compositions have been tested, from the zirconia-rich to the alumina-rich side [89, 90]. Major ceramic companies are developing such materials. A composite material processed with 80% tetragonal zirconia poly crystals (Zr-TZP) and 20% alumina (Al₂O₃) is reported to have outstanding mechanical and tribological properties. The alumina-toughened zirconia (ATZ, rather than ZTA) Bio-Hip^s, developed by Metoxit AG (Switzerland), has a bending strength of up to 2000 MPa, which is extremely high for a ceramic material. CeramtecAG (Germany) recently developed BIOLOX^sdelta, which consists of approximately 75% aluminum oxide, the basis for hardness and wear resistance, and approximately 25% zirconium oxide, for improved mechanical properties. A strength higher than 1150 MPa is reported, associated with a toughness of 8.5 MPa.m^1/2. This commercially available product belongs to the conventional ZTA composition. The addition of alumina to zirconia clearly hinders aging or at least reduces its kinetics. The material does not lose strength after repeated steam sterilization cycles. However, Chevalier et al. warn researchers to be careful. As an example, they showed that that aging could be significant in a 3Y-TZP – alumina composite above 16 vol% zirconia [91]. This critical content is related to the percolation threshold above which a continuous path of zirconia grains allowed transformation to proceed. The conclusion is yttria as a stabilizer still poses problems; however, there are several strategies that can be pursued to avoid such problems.

Glass-Infiltrated Ceramics  – These materials are based on partially sintered alumina, zirconia, and spinel. These materials are commercially known as In-Ceram Spinell®, In-Ceram Alumina®, and In-Ceram Zirconia®. They have surface pores, which are infiltrated with lanthanum aluminosilicate glass. Lanthanum additions decrease the viscosity of silicate melts, which assists in infiltration and simultaneously improves the strength of the ceramic core. The mechanical properties of In-Ceram® materials are better than all other restorations other than Y-TZPs. From the esthetics point of view, In-Ceram® materials are opaque and they can be tinted easily.

14.4.3 Processing of All-Ceramic Restorations

A crown with the normal size and shape of a tooth is very useful, when a tooth is broken down to a limit when filling no longer solves the problem. Traditional crowns have a porcelain-bonded enclosure filled with gold. They are sturdy enough to withstand about 200 pounds of pressure during use. Currently, second-generation materials are being used which are stronger and more durable than the original porcelain type materials with better wear, and more translucency to match the natural color of the teeth. All-ceramic restorations (e.g., crowns, inlays, onlays and veneers) are becoming quite popular when esthetics is important. A bridge is a complex version of a crown replacing one or more natural missing teeth, thereby “bridging” the space between two teeth. Bridges are cemented into place on the “abutment” teeth – the surrounding teeth on either side of the space or span. Unlike removable partial dentures, bridges cannot be taken out of the mouth by the patient. There are several criteria for selecting restorative systems and processes. Esthetics, strength, marginal fit, biocompatibility, cost and ease of fabrication, and patient comfort are among these criteria. A process and a relevant material should be considered most versatile when the combination can successfully produce a bridge.

Most of the dental prosthetics that are produced manually require considerable technical skills and consume a substantial amount of time. Machining has become a viable option for most of the dental restorations [92, 93]. Computer-assisted design/computer-assisted manufacturing (CAD/CAM) has become a viable option for making all ceramic restorations. Of late, CAD/CAM is referred to as the subtractive mode of rapid-prototyping technology whereby blocks of materials are machined to precise shapes. Commercially available systems include CEREC (Sirona), Procera (Nobel Biocare), and LAVA (3 M ESPE). CEREC mechanics are well suited for machining glass ceramics, In-ceram alumina, spinel, and zirconia [94]. While LAVA machines can also machine Y-TZP materials. The prepared teeth are coated with an imaging liquid, which can hold imaging powder such as TiO₂ on the teeth. Using an infrared camera, an optical impression is captured and suitably modified for the proper fit. The data are transferred to a milling machine to mill a block of ceramic. After milling, the sprue is removed from the restoration. The main advantage of these machines is the elimination of a temporary restoration, resulting in a significant reduction of time and cost. They also reduce the human error in processing. The difference between the CEREC and LAVA machines is that the CEREC machines can fabricate one crown while the LAVA machine can produce bridges. This can be achieved because of the use of Y-TZP of high strength and toughness. An alternative to the LAVA system is the Cercon® Zirconia system, which requires sending an impression to a laboratory for machining a pre-sintered porous Y-TZP. Once the machining is carried out, the restoration is sintered to full density taking into account the shrinkage.

The use of all ceramic materials requires a tooth preparation with a rounded shoulder or chamfer finish line. The amount of tooth reduction varies from system to system [95]. For example, the manufacturer of Empress 2 ® recommends 0.8 mm, while Procera ® and In-Ceram ® recommend 0.5 mm. Cementation with composite resin cements improves the success rate of Empress 1 and 2. Glass-ionomer cements are popular because they can be expanded and create stress, which will eventually lead to failure. Hybrid cements such as copolymers (poly-acid modified resin cements and the resin-modified glass-ionomer cements (GIC) were developed as smart filling materials. However, biocompatibility remains an issue [96].

14.4.4 Selection Guide for All-Ceramic Restorations

In selecting materials, higher strength and fracture toughness should be used for canine and posterior regions of mouth. Besides, Weibull modulus is an important factor for successful restorations. If the scattering of tensile strength is high, Weibull modulus is low; therefore, reliability might be a concern.

The strength of castable leucite reinforced glass ceramics is around 130 MPa, and they are best suited for anterior crowns, inlays, onlays, and veneers. In-Ceram zirconia has a significantly higher strength and fracture toughness and can be used for anterior and posterior single unit crowns and anterior bridges. Lava form® (Y-TZP) has the highest strength and toughness. Table 14.7 shows selection guide for the use of all-ceramic restorations.

Table 14.7 Selection guide for all-ceramic restorations

14.4.5 Clinical Failure of All-Ceramic Crowns

Despite the wide variety of all-ceramic restorative materials available today, they are still not highly recommended by dentist practitioner [96]. Usually cementation or internal surface flaws initiate most of the clinical failures. As an example, Dicor ceramic crowns luted with Zn-phosphate cement were reported to have poorer success rates than crowns cemented and bonded with composite resin cements. Internal surfaces have the highest tensile stresses and critical flaws. Etching and polymer coating of tensile surfaces can strengthen the strength of these ceramic materials by blunting the crack tip or by lessening the stress corrosion by barrier coatings [97,98,99].

14.4.6 Bioactive Glasses

Bioactive glasses are widely used in musculoskeletal applications, a smaller subgroup of which is dentistry. One of the important systems of bioactive glasses is derived from Na₂O–CaO–P₂O₅–SiO₂ ternary system with other substitutions. The system is commercially known as Bioglass®, which was developed by Hench et al. at the University of Florida [100, 101]. An important feature of these glasses is a lower SiO₂ content as opposed to the traditional soda-lime silica glass and much higher content of Na₂O and CaO. These compositions make the surfaces highly reactive when exposed to the body fluid. The most commercially important composition is known as 45S5 containing 45 mol% SiO₂.

Tomsia and colleagues modified the original compositions so that they are compatible to metallic implants in terms of thermal expansion coefficients [102,103,104,105]. Specifically, Na₂O and CaO were partially substituted by K₂O and MgO, respectively. This results in thermal expansion coefficients compatible to implants. The processing of such glasses can be performed by the traditional glass melting process or by the sol-gel techniques.

As mentioned before, monolithic net-shaped implants have been in use in dentistry. Specifically, Endosseous Ridge Maintenance Implant (ERMI) was developed by Stanley et al. [106]. Their 10-year survival rate is greater than those implants that do not osseointegrate. Bioglass® particulates have been placed around teeth with periodontal disease [107]. Osseointegration takes place at a rapid rate around those glass particles. This helps in stabilizing the boundary between tooth and the periodontal membrane and the tooth is saved.

The other related material is actually a glass ceramic based on the CaO–P₂O₅–SiO₂ system. This material consists of crystallites of apatite Ca₁₀(PO₄)₆ (OHF₂) and β-wollastonite (CaO–SiO₂) with a residual CaO–SiO₂ glassy phase. These materials are referred to as A/W glass ceramics and were mentioned by Yamamuro and Kokubo [108,109,110]. The presence of crystallites enhances the mechanical properties (e.g., fracture toughness, strength as well as subcritical crack growth velocity). The mechanism of bioactivity of such glass ceramics is similar to the Bioglass®. By the same token, these can also be used in the cases where Bioglass® is used.

14.5 Polymers for Dental Applications

Polymeric materials with improved durability find important applications as dental materials. Denture bases and artificial teeth are two major areas where polymeric materials are used. Besides these, they are also being used in various applications such as crown and bridge facings, inlay patterns, implants, impressions, dies, temporary crowns, endodontic fillings, athletic mouth protectors, and orthodontic space maintainers, etc. [111].

14.5.1 Dentures

Dentures are usually used if an entire arch of teeth is missing. If there are a few teeth left, partial denture is needed. Generally, dentures have an acrylic base molding. The quality of the bonding of denture base polymers to synthetic polymer teeth is dependent on the temperature during processing. About 98% of the denture materials are constructed from methyl methacrylate polymers or copolymers. Other polymers include vinyl acrylic, polystyrene, epoxy, nylon, vinyl styrene, polycarbonate, polyurethane, silicones, rubber-reinforced acrylics, and butadiene-reinforced acrylics, etc. Systems which are based on dimethacrylate monomers are commonly used for crowns and bridges, and now also for dentures.

The most important polymer system used for prosthodontic applications is based on polymethyl methacrylate (PMMA) powder and a monomer mixture (liquid) of methyl methacrylate and a cross-linking dimethacrylate monomer such as ethylene glycol dimethacrylate (EGDMA). The process of polymerization is crucial for the properties of the prosthodontic devices. The mechanical properties including craze resistance are dependent on the processing procedure and on the quality of cross-linking agent. At an ambient temperature, the material is brittle and transparent. At temperatures above the transition temperature, the material tends to soften and becomes viscoelastic.

The presence of surface cracks/pores and internal faults usually affect the tensile strength of the material. The failure of the denture material involves either impact or fatigue. The impact failure is due to rapid stress of the material, and fatigue failure is due to cyclic flexing of the material generating cracks.

14.5.2 Dental Cements

Dental cements are used for fixing dental restorations such as crowns, orthodontic bands, serving as bases under any restorations, etc. They must exhibit sufficient viscosity to flow through the interfaces between hard tissue and the restoration [12]. Properties that are important include strength and stiffness, resistance to dissolution, and biocompatibility with pulpal tissue [111]. Further, they should be set within 25–50 μm thickness, and should release fluoride to prevent caries. The five most popular systems include zinc phosphate, zinc oxide eugenol (ZOE), polycarboxylate (with zinc powder), glass-ionomer, and compomer cements.

Zinc Phosphate and ZOE – Zinc phosphate serves as an universal luting agent, which means that it has the right rheological properties to flow between two surfaces with the ability to solidify very fast [12]. It has good handling characteristics, and a proven longevity in the oral cavity when used in conjunction with well-designed and well-fitted restorations. The powder constituents for both zinc phosphate and ZOE are zinc oxide with minor additions of magnesium oxide. The main liquid component is either phosphoric acid or eugenol oil. Zinc phosphate is typically used for permanent cementation. The accepted mechanism for cementation is the initial attack of oxides by phosphoric acid, resulting in an aluminophosphate gel. The problem with zinc phosphate cement is that it is highly acidic and it may cause damage to the pulp. On the other hand, the advantage of ZOE is its biocompatibility, eugenol having a pH of 7 and hence benign to the pulp. Zinc oxide is hydrolyzed initially and undergoes chelation subsequently. ZOE can be used for either permanent or temporary fixation. There are several variations in the compositions of ZOE based on their service demands. These cements are used as a base where a significant amount of dentin has been removed and pulp tissues are approximately 0.5–2 mm. Rather than placing a restoration near pulp, cements are used to protect the pulp.

Polycarboxylic Acid Cements – In polycarboxylic cements, the liquid component is polyacrylic acid , which is then mixed with zinc oxide [112]. Bonding to dentin can occur via chelation of carboxyl groups to calcium ions. The advantages include shorter gelling times (about half of zinc phosphates) and better bonding strength with the tooth structure (with respect to other zinc phosphates) [113].

Glass-Ionomer Cements  – These use silicate glass powder instead of zinc oxide powder and an aqueous solution of polyacrylic acid (or a copolymer of itaconic acid, or maleic acid) [114, 115]. They bond to tooth structure well and release fluoride. Typical composites of glass particles are 20–36 wt% SiO₂, 15–40 wt% Al₂O₃, 0–40 wt% CaO, 0–35 wt% CaF₂, 0–5 wt% Na₃AlF₆, 0–6 wt% AlF₃, 0–10 wt% AlP. When glass particles are mixed with polyacrylic acid, hydrogen ions are removed from carboxyl groups of acid solution. Simultaneously glass particles undergo reaction, where Al³⁺, Ca²⁺, Na⁺, F⁻, and PO₄ ³− are released into solution to form a silica-rich gel on the particle surfaces. In order to use glass-ionomer cements, the dentin and enamel are acid etched with polyacrylic acid, then rinsed. Bonding to dentin occurs by both ionic and micromechanical forces. Glass-ionomers have the advantage of chemically bonding directly to enamel and dentin. There is an ionic bond formed between the calcium of the tooth and the set material. As enamel is richer in calcium, the bond to enamel is stronger than the bond to dentin. The bond strength of glass ionomer is much lower than that of bonding agent/composite systems, but in some situations, high bond strength is not the priority. In terms of toughness, they cannot withstand high stress concentrations that promote crack propagation . Besides, slow maturing contributes to low initial strength. Sometimes, glass-ionomers are reinforced to “cermets” (silver or silver amalgam powder mixed into original powder or pure metallic silver fused to glass powder) for increased strength, more radio opacity, and increased abrasion resistance [116]. Another type of glass-ionomer cement is resin modified (hybrid ionomer), which overcomes the problem associated with low initial strength. Polymerizable functional groups can be added to impart rapid curing when activated by light or chemicals. This overcomes the problem associated with low early strength and moisture sensitivity that occurs during slow acid base setting reactions.

Compomer – Compomers have the fluoride-releasing ability of glass-ionomers and the durability of composites. These are typically one component systems consisting of fillers of sodium fluoride and polyacid-modified monomers. The setting reaction is triggered by photopolymerization. The mixture does not contain water and hence is not self-adhesive like the glass ionomer cements. A separate dentin bonding agent is therefore needed.

Resin Cements – Resin cements are used for attaching orthodontic brackets and fixed prostheses to tooth structure following the application of either enamel or dentin bonding agent. These cements are insoluble in any oral fluids. Although resin cements have lower elastic modulus compared with zinc phosphate, fracture toughness is usually superior. Resin cements bond well with the dentin and make a strong attachment with the enamel.

14.5.3 Composite Dental Materials

Dental composites are becoming popular because of esthetic reasons and are used to rebuild or repair the teeth. The compositions consist of three parts – a resin, filler particles, and a coupling agent (typically a silane to bond filler with the resin). Composite resins have good esthetic properties and strength and are used in front teeth. The resin that is of key importance in dental composites is an aromatic dimethacrylate monomer, bisphenol A-glycidyl methacrylate (BIS-GMA) [117]. Polymer shrinkage, thermal expansion, water sorption, tensile strength, wear resistance, and modulus of elasticity are some of the properties that can affect clinical performance. In order to enhance further the longevity, filler particles such as quartz, colloidal silica, or silicate glasses containing Sr or Ba are incorporated into the resin matrix. These tooth-colored filling materials have reached a high degree of sophistication since their appearance on the dental scene in the early 1960s. A modern dental composite consists of a paste created by combining a mixture of dimethacrylate monomers and cross-linking agents with up to 80 wt% of silane-coated, ceramic particles (the filler), whose sizes range from 0.04 to 4 μm. This composite paste is packed into a dental cavity and the dentist exposes it for about 30s to intense visible blue light. The light activates a chemical initiator (camphorquinone, CQ) within the composite and the resins undergo free radical addition polymerization via their vinyl groups, turning the paste into a durable, solid filling. Fillers are placed in dental composites to reduce shrinkage upon curing. Physical properties of composite are improved by fillers; however, composite characteristics change based on filler material, surface, size, load, shape, surface modifiers, optical index, filler load, and size distribution. Accordingly, the composites are classified as traditional (containing large fillers), hybrids (containing different particle sizes), flowable hybrids (containing different particle size with lower volume fraction of fillers), and homogeneous microfill (containing ultrafine/nanocrystalline fillers). Fillers less than 1 μm do not produce a rough-appearing surface with aging. The nanoparticles fill between all other particles to further reduce shrinkage.

Composite fillings are less durable than amalgam for the large filling size. However, it is comparable for small-to-average size fillings. Usually composite fillings are quite difficult to place especially for a back tooth, and therefore, they are more expensive than amalgam filling. Composites are used in areas where fillings are relatively small and there is plenty of tooth structure for support. They have similar strengths to amalgam but they tend to wear away more rapidly. Composite fillings are more natural looking, require less tooth reduction to place, and are bonded in place for a better seal. In terms of biocompatibility, composites are not totally biocompatible. Sometimes iron oxide, aluminum oxide, barium, and other materials leach out.

Polymerization shrinkage is one of the major disadvantages of composite materials. As the resin matrix polymerizes, the organized polymer molecules occupy less than its disorganized molecules. During cures, it shrinks considerably. The degree of shrinkage is determined by the percentage of filler in the composite (most hybrids are 60–65% filler by volume).

14.5.4 PEEK (Polyetheretherketone) as a Dental Material

Of late, there has been a surge in the application of polyetheretherketone (PEEK) in dentistry because of its interesting properties as compared to those of metals. An important application in dentistry is in dental implants [118, 119]. Most notably, the Young’s modulus of PEEK is 3-4 GPa, which is similar in range with that of bone. The material has a low density (1.38–1.4 g/cm³). The comparable Young’s modulus values with those of bone make implants less prone to stress shielding. PEEK is resistant to chemical, radiation, and thermal energy exposures. Finally, it is biocompatible, causing neither mutagenic effects nor clinically significant inflammation. Taken together, these materials are quite viable as alternative to metallic implant materials.

PEEK is a member of the polyaryletherketone (PAEK) family. Because of the reasons presented in the above, PAEK materials have a long history of being used in fabricating orthopedic, trauma, and spinal implants. PEEK has emerged as the leading high-performance thermoplastic candidate for replacing metallic implants, especially in orthopedic trauma [120, 121]. According to the recent literature, there is a growing interest in using PEEK in dentistry including implantology [122, 123]. The material has been popularized after successful demonstration of the nucleophilic synthesis route by ICI in the UK paved the way for commercial production of PEEK [124]. It is now well established that commercially available medical grade materials are easily amenable to standard polymer manufacturing processes such as injection molding, extrusion, and compression molding.

The PAEK group of polymers consists of aromatic rings joined by ether or ketone linkages with specific polymer name describing the number and order of linkages; these highly aromatic polymers are linear, with little or no branching [125]. These structural features explain their impressive inertness. Chain configurations are controlled by the steric hindrance resulting from the aromatic groups and the bond angles of ether and ketone linkages. The average inter-ring linkage angle is reported to be between 124° and 126.5° for PEEK [125, 126]. PEEK polymers are typically semicrystalline, more specifically PEEK can be 40%, even though 30–50% is more typical [127]. The result is a two-phase morphology with crystallites embedded within amorphous regions. According to X-ray results, PEEK is orthorhombic, belonging to the space group Pcbn. The unit cell dimensions are a = 0.775 nm, b = 0.586 nm, and c = 0.1 nm [126]. Assuming these dimensions, the calculated density is 1.4 gm/cm³ [118, 126].

Processing of PEEK is a challenging task as opposed to traditional polymeric materials. During cooling from the melt, a quenching step can produce amorphous materials, which is transparent and lacks the chemical and thermal durability. Thus, semicrystalline microstructure is preferred. During crystallization, the PAEK polymers align and fold into lamellae, excepting when nucleants are present. Traditional characterization techniques as wide-angled X-rays or simply X-ray a diffraction (XRD), fourier transform infra-red (FTIR) spectroscopy, thermal analysis such as differential scanning calorimetry (DSC), scanning electron microscopy (SEM) are important in the characterization of the substrates. Of these, DSC is expected to be very important to study if any thermal degradation occurred during the coating process. XRD peaks at 18.5° and 23.5° are the characteristic peaks of PEEK [100]. In the case of FTIR data, the ratio between absorption bands at 1305 and 1280 cm⁻¹ also show crystallinity. The glass transition temperature, Tg, is between 125 °C and 165 °C, whereas the melting points are between 260 °C and 360 °C [128]. The glass transition temperature has a profound effect of the material properties [128]. Since Tg is closely related to processing conditions, it is important to perform such characterizations just to ensure that the polymeric substrate has not deteriorated during processing.

As compared to metallic implants, the major drawback of PEEK is its lack of osseointegrability. To enhance osseointegration of PEEK implants, one strategy is to modify their surfaces using some kind of physical or chemical treatments. The commonly used physical treatments include exposure to plasmas. Also, the recent decade has witnessed the development of various kinds of multi-functional coatings on PEEK. As bacterial colonization on the dental implants remains a grievous problem, a significant amount of research has been focused on developing bioactive and antibacterial coatings on PEEK and other implant materials [129,130,131,132]. Chemical treatments include sulfonation treatments or exposure of surfaces to electromagnetic energies. Better yet, it is possible to make composites following polymer-processing techniques by including inorganic particulates. Table 14.8 presents the state-of-the art of this field.

Table 14.8 Bioactivity of PEEK composites

14.6 Closure

Current dental restorative materials can be used effectively for restoring teeth for functional or esthetic reasons. Diverse materials such as metals, ceramics, polymers, and composites are routinely used. Important properties include reliable mechanical properties, corrosion/wear resistance, biocompatibility, and physical appearance. No single material has been found that meets all the requirements. Furthermore, virtually all restorative materials have components with potential health risks. However, there is no scientific evidence that currently used restorative materials cause significant side effects. Available data do not justify discontinuing the use of any currently available dental restorative materials or recommending their replacement. The research is going along many different avenues. It is hoped that in the future, there will be substantial volume of research in this field.