Oshida Y. Bioscience and Bioengineering of Titanium Materials

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Implant Application 253

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Chapter 9

Other Applications

9.1. Denture Bases 257

9.2. Crowns and Bridges 259

9.3. Clasps 265

9.4. Posts and Cores 266

9.5. Titanium Fiber and Titanium Oxide Powder as Reinforcement 267

for Bone Cement

9.6. Sealer Material 268

9.7. Shape-Memory Dental Implant 268

9.8. Orthodontic Appliances 269

9.9. Endodontic Files and Reamers 271

9.10. Clamps and Staples 273

References 274

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257

Chapter 9

Other Applications

This section will describe a variety of titanium applications other than the implants

discussed in the previous chapter. Some applications are realized owing to

advanced and modified manufacturing technologies, and other applications are

developed and promoted due to material’s uniqueness, such as superelasticity

and/or SME associated with NiTi alloys.

9.1. DENTURE BASES

Denture materials are materials, which require precise fabrication (or these pros-

theses should be manufactured in a near-net-shape form). Traditionally, various

types of polymers have been used, including polymethylmethacrylate (PMMA),

polycarbonate (PC), polystylene, and polysulfone. However, any polymeric resins

have drawbacks. These include weakening of mechanical properties due to water-

sorption, which also causes a dimensional instability, and tissue irritation due to

unavoidable residual unpolymerized monomer.

CpTi and Ti-based alloys are generally used as fixtures for dental implantation

[9-1, 9-2]. When these metals are used as restoratives for anterior teeth, the metal

prostheses are generally covered with esthetic materials, such as porcelain or resin.

Veneering resin is more popular than porcelain because it is less expensive and

more resistant to abrasion [9-3]. When constructing removable dentures, implant

supported prostheses and most especially resin-bonded bridges, the interfacial

bond between metal and acrylic resin has been wholly reliable or durable [9-4].

Numerous investigations have been carried out to measure metal/acrylic resin

bond strengths between materials that have surface or chemical modifications that

are reported to enhance retention [9-49-6]. These have included surface treat-

ments of the alloys by macro-, micromechanical and chemical methods, or by the

addition of chemical bonding agents to the acrylic [9-6, 9-7]. PMMA resins have

a linear thermal coefficient of expansion of 8110

6

m/m·

C, whereas noble and

base metals range from 7 to 1910

6

m/m·

C [98]. The lack of adhesion and dif-

ferences in the thermal expansion of the materials places severe demands on the

metal/acrylic resin surfaces. Ban [9-1] prepared CpTi (grade 1), Ti-6Al-4V, exper-

imental -Ti alloy (54Ti-29Nb-13Ta-4.6Zr), and 12Au-20Pd-52Ag-15Cu-1Zn

alloy plates which were sand-blasted with alumina Al

powder (100250 m),

water-washed and dried. CpTi plates were further oxidized in-air at 600

C for 1 h.

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They were also treated in 5 M LiOH, 5 M KOH, and 5 M NaOH at 25 or 60

C for

20 h, followed by in-air oxidation at 600

C for 1 h. The bonding strengths of such

treated CpTi plates were measured using a pull-shear bonding method after

immersion in physiological saline solution at 37

C for 24 h. It was reported that (i)

the bonds of the standard alkaline-treated CpTi and two Ti alloys to resin were

1.51.9 times stronger than those of sand-blasted specimens, but no significant

effects of the alkaline treatment were observed on 12Au-20Pd-52Ag-15Cu-1Zn

alloy, (ii) the greatest bonding strengths were found for CpTi treated with NaOH

and KOH and the heated at 600

C, and (iii) alkaline treatment is a simple, effec-

tive surface modification of Ti that improves bonding to veneering resin [9-1].

The recognition that a chemical linkage between acrylic resin and metal would

lead to an enhanced bond led to the development and use of an intermediate layer

based upon silicon oxide. The so-called “SiO-C” layer was introduced as an inter-

mediate layer between the metal substrate and polymer as a means for enhancing

the strength and durability of the bond [9-9]. Ti-6Al-4V, after bonding (n  3), was

stored at 37

C in distilled water for 7 days, followed by thermocyclings between 5

and 55

C for 500 cycles (for 21 days), and then was subjected to the 4-point bend-

ing at a crosshead speed of 0.5 mm/m. The bond strength   F(L  ᐍ)/R

where  is in N/m

, F the load at failure (N), L the distance between specimen sup-

port (m)  0.024 m, ᐍ  a distance between loading bars (m)  0.008 m, and R

is a radius of specimen (m)  0.0035 m. It was concluded that the most durable

surface treatment was the silicoating process after sand-blasting of the metal

surface with 250 m alumina for heat-cured acrylic resin materials [9-10].

Denture bases are now mainly manufactured by a casting process of Co-Cr or Ni-

Cr alloys. A novel superplastic fabrication process for dentures, as well as implant

systems, has been developed [9-11]. Ti-6Al-4V alloy is known as a biocompatible

metal, and with both high strength and low density, it seems to be the best material

for denture bases. The Ti-6Al-4V also has superplastic properties at high tempera-

ture, forming very simple processes by argon pressure (10 kg/cm

) at 800900

Superplastically formed (SPFed) denture bases of Ti-6Al-4V showed some merits

compared with conventional cast or cold-pressed denture base plates, namely, good

fitness, light weight, and low producing cost using a smaller number of manufac-

turing processes [9-12]. A SPFed Ti-6Al-4V removable denture framework was

fabricated. Clinical application in short-term follow-up periods from 6 months to 3

years was evaluated. It was reported that (i) the dentures functioned well and did

not cause any major clinical difficulties, and (ii) the patients have expressed satis-

faction with the dentures at regular recall appointments [9-13].

Srimaneepong et al. [9-14] cast Ti-6Al-7Nb alloy into three differently

designed, removable partial denture frameworks: palatal strap, anterior–posterior

bar, and horseshoe-shaped bar. The vertical displacement and local strain of

258 Bioscience and Bioengineering of Titanium Materials

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Ti-6Al-7Nb alloy frameworks were investigated to compare against those of Co-Cr

alloy frameworks. A vertical loading force of 19.6 N was applied at two locations,

10 and 20 mm, from the distal end of the frameworks. It was found that (i) although

higher vertical displacement and local strain were observed for Ti-6Al-7Nb alloy

frameworks than those for Co-Cr alloy frameworks, the palatal strap framework

appeared to show the least deformation, (ii) the strain at 10 mm location was higher

than that at 20 mm location for anterior–posterior bar and horseshoe-shaped frame-

works, and (iii) the palatal strap design was evaluated to be a suitable design for the

removable denture framework with Ti-6Al-7Nb alloy [9-14].

9.2. CROWNS AND BRIDGES

Dental porcelains provide dentistry with biologically acceptable and aesthetic

(tooth-colored restoration) methods of repairing diseased and fractured teeth [9-

15]. Modern dental porcelains are pigmented glass-crystal composites that were

evolved from traditional ternary whiteware compositions, namely clay-feldspar-

quartz. Most of the modern dental porcelains used for the fabrication of crowns and

bridges require only short, low-temperature sintering cycles. Porcelains have been

fluxed so that their thermal expansions match precious or non-precious alloys. Such

alloys are manufactured into frameworks for the crowns or bridges, and the porce-

lains are bonded to them. The metal crown or bridge is faced or veneered with

porcelain. The bond between precious alloys and the porcelain is more reliable than

the bond between non-precious alloys and the porcelain. The presence (by alloying

in the original coping materials or coating on the coping materials) of small

amounts of tin (Sn) or indium (In) in the precious coping alloys makes the chemi-

cal bonding possible in these cases. It is generally believed that oxygen in InO and

SnO shares with oxygen in the silica component in porcelain, resulting in enhanced

bonding strengths. All these porcelains have been opacified as to cover their dark

metal-oxide background. They are highly reflective and are inferior aesthetically for

use in anterior teeth. Porcelains formulated for metal-ceramic prostheses have a

tendency to densify during long and successive firing cycles [9-15]. porcelain-

fused-to-metal (PFM) is sometimes called porcelain-bonded-to-metal (PBM),

porcelain-to-metal (PTM), ceramometal, or metal–ceramics. According to fusing

temperature, porcelain can be classified into (1) high fusing (1300

C) for denture

tooth fabrication, (2) medium fusing (1100–1300

C) for denture tooth fabrication,

(3) low fusing (8501100

C) for crown/bridge construction, and (4) ultra-low

fusing (<850

C) for crown/bridge construction. As for metallic coping materials,

they are also classified into (1) high noble (Au-Pt-Pd, Au-Pd-Ag, Au-Pd), (2) noble

(Pd-Au, Pd-Au-Ag, Pd-Ag, Pd-Cu, Pd-Co, Pd-Ga-Ag), and (3) base metal (CpTi,

Ti-Al-V, Ni-Cr-Mo-Be, Ni-Cr-Mo, Co-Cr-Mo, Co-Cr-W).

Other Applications 259

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There are several requirements and characteristics of metals applicable to the

metal-ceramic systems as follows.

(1) The metal should have a high melting point so that it does not melt during the

porcelain application procedure.

(2) The metal should have a similar coefficient of thermal expansion as those val-

ues in porcelain. The coefficients of thermal expansion (linear),  (in  10

6

m/m·

C, or simply in ppm), for major materials are reported as follows: con-

ventional high-fusing porcelain (13.715.5 ppm), conventional low-fusing

porcelain (14.115.1 ppm), low fusing porcelain for Ti veneering (7.7–8.9

ppm), Au-Pd (13.9 ppm), high Pd (14.7 ppm), Ni-Cr base metal (13.9 ppm),

CpTi (9.7 ppm) [9-16].

However, a slight extent of thermal mismatching is

favored by 

METAL

> 

PORCELAIN

, so that the porcelain will be placed under the

compression on cooling.

(3) The metal should have excellent castability (however, casting is not only a

fabrication method, but it should also include CAD/CAM or superplastic forming

 SPF).

(4) Both materials should have similar values of modulus of elasticity to establish

the mechanical compatibility [9-17].

Moreover, since little or no residual stress due to thermal mismatch should exist

in the final Ti/porcelain-bond interface, the significant discrepancies already

noted in their thermal coefficients of expansion will have to be modified to more

closely match. It appears that an interfacial oxide layer, some 100–1000 nm

(10–100Å) thick, forms during firing, and the thicker this layer becomes, the

weaker the bonding strength between the porcelain and the titanium [9-18].

To have a reasonable function of Ti/porcelain system, the rigid bonding is crucial

in terms of in-service longevity. There are several theories to postulate bonding

mechanism of porcelain.

(1) Chemical bonding. A chemical bonding, existing through the formation of a

tin (or indium) oxide layer during porcelain firing and that the oxide formed

provides the mechanism of bonding between the porcelain and the metal sub-

strate. Watanabe et al. [9-19] characterized the interface reaction between

porcelain and pure titanium during porcelain firing, using an electron–probe

microanalyzer (EPMA), X-ray diffraction method, and Fourier transform

260 Bioscience and Bioengineering of Titanium Materials

When the volumetric expansion, , needs to be calculated, it follows as:   (1 )

1

 (1  33



)  1{ 3, since 

(in order of 10

12

) and 

(in order of 10

18

) are small

enough to be neglected.

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