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The goal of modern dentistry is restoring the patient to normal condition

of function, esthetic, speech, and general health (Misch, 2003).

Teeth are lost due to many reasons as: tooth decay, root canal failure,

periodontal disease, trauma to the mouth (tooth injury), excessive wear and

congenital defects (Allisn et al., 2009).

Many complications occurred due to loss of the teeth as proximal drafting

of neighboring teeth and over eruption of opposing teeth that lead to opening

of contact of teeth, increase of caries incidence, periodontal problem and

esthetic complications (Misch, 2003). In addition to psychological problems

as depression and acute crises of self confidence will occur especially in

young age when one of anterior teeth loosed (Fiske and Davis, 2008).

Traditional solutions for tooth missing are fixed and\or removable partial

dentures:

Removable partial denture has many disadvantages; as its instability and ill

retention as natural tooth, beside patient's discomfort and it attaches to natural

teeth by visible clasps and hooks, cause stress on the natural teeth which can

loose them, and promote tooth decay (Aquilino and Shugars, 2001).

In spite of fixed partial denture more stable and retentive than removable

one, there're some disadvantages as hazard on neighboring teeth occurred by

the reduction. Additionally to its durability and stability depend on status &

healthy of abutment teeth (Shugars et al., 2003).

All of above solutions have one drawback as: alveolar bone loss. The

alveolar bone that supports teeth will be resorped after loss of the teeth. Bone

resorption will occur even with the best bridges, partial or complete dentures.

Bone loss also causes changes in facial features that can make you look older

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sooner by causing: a deepening of the groove between the nose and the

corners of the mouth, pronounced, forward jutting jaw, beside shortening of

the distance between the chin and nose and sagging of the facial muscles and

unsightly jaws (Quirynen et al., 2007).

To overcome this drawback, dental implant is indicated. It depends on the

jawbone in its retention and stability and has many advantages as; good

esthetic, improved phonetics, increased stability and retention, in addition it

reduced bone resorption (especially around implant) (Priest and Priest,

2004).

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Dental Implant can be defined as: 'a permanent device that is biocompatible

and bio functional, inserted in the bone of the jaw to replace missed tooth and

provide retention and support for fixed or removable prosthesis' (Mortilla

and Lynn, 2009).

Osseo integration implies a direct connection between a vital bone and

screw shaped titanium implant with defined finish and geometry-fixture

(Mortilla and Lynn, 2009).

II.1 History of dental implant:

Firstly from 3000 years ago: the first copper stud was implanted into

Egyptian mouth (Arbree, 2005). This stud was also observed in mandible of

Mayan mummy during excavating Mayan burial sites in Honduras in 1931.

Archaeologists found a fragment of mandible of Mayan origin, dating from

about 600 AD, this mandible had three tooth shaped pieces of shell placed

into the sockets of three missing lower incisor teeth and compact bone was

formed around the implants which led archaeologists to conclude that the

implants were placed during life (Crespi et al., 2008).

The transplantation of natural teeth which extracted from poor individual

to upper social classes was common during eighteenth & nineteenths

centuries especially in western civilization (El-Askary, 2008).

In 1809, Magiolo (Nancy France) forcing a metallic tooth root made of 18k

gold into socket of extracted tooth which usually followed by gingival

inflammation and severe pain . While Dupont reimplanted the extracted teeth

with endodontically treated into empty socket (Bruijn et al., 2001).

Then Edmunds& Harris in 1886 implanted a platinum post & porcelain

crown into artificially created socket in the alveolar bone, the post was

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covered by bone and some implants were still in function after 27 years of its

initial placement (El-Askary, 2008).

Adams in 1937 presented first submerged threaded cylindrical implant, this

design included smooth gingival collar and healing cap (Crespi et al., 2008).

In 1952 the Swedish orthopedic surgeon, P I Brånemark, was studied the

bone healing and regeneration, and observed that the bone was effectively

adhered to the titanium and grown in contact with it. Brånemark carried out

many further studies into this phenomenon, using both animal and human

subjects, which all confirmed this unique property of titanium (El-Askary,

2008).

Brånemark decided that the mouth was more accessible for continued

clinical observations and the high rate of edentulism in the general population

offered more subjects for widespread study. He termed the clinically

observed adherence of bone with titanium as ‘Osseo integration’. In 1965

Brånemark placed his first titanium dental implant into a human volunteer, a

Swedish female named Gösta Larsson (Arbree et al., 2005).

Meanwhile an Italian medical doctor called Stefano Melchiade Tramonte,

concluded that titanium could be used for dental restorations (implants) and

after designing a titanium screw to support his own dental prosthesis, started

to use it on many patients in his clinic in 1959. The good results of his

clinical studies on humans were published in 1966 (Crespi, et al., 2008).

II.2 Types of implants (as general)

Autoplastic implants e.g.: implant from- into the same individual.

Homoplastic e.g.: implant from same species.

Heteroplastic e.g.: implant from different species.

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Alloplastic e.g.: implant from non living material (Arbree et al., 2005).

II.3 Types of dental implants:

End osseous, Subperiosteal and Tran osseous Implants.

II.3.1 End osseous implants: which are surgically inserted into the jawbone

(root form, blade form, ramus blade and ramus frame). (Stuart and Green,

2004)

II.3.1.a Root form implants:

Root form implants are the closest shape and size to the natural tooth root.

They are commonly used in wide, deep bone (more than 8mm in heights,

5,25mm buccolingually and 6,5mm mesiodistally) to provide a base for

replacement of one, several or a complete arch of teeth to support fixed, fixed

detachable over denture, and single tooth crown. (Abouzga and Games,

2002).

There are variations of the root form implant dwell on their shape. Some are

screw-shaped, others are cylindrical, or even cone-shaped or any combination

there of (Park et al., 2008).

II.3.1.b Blade form implants:

Blade implants are not used too frequently any more, however they find an

application in areas where the residual bone ridge of the jaw is either too thin

to place conventional root form implants (more than 8mm height, 3mm

buccolingually and 10mm mesiodistally) (Cranin, 2001).

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II.3.1.c Ramus frame implants:

One of end osseous implants, although their appearance are not look like

end osseous implants. These implants are designed for the edentulous 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. The part of the implant

that is visible in the mouth after the implant is placed looks similar to that of

the Subperiosteal implant (Abouzga and Games, 2002).

It indicated in a severely resorbed, edentulous lower jaw bone (more than

6mm height and 3mm buccolingually) to give support to only over denture,

also the advantage that comes with this type of implant is a tripodial

stabilization of the lower jaw from fracturing (Stuart and Green, 2004).

II.3.2 Subperiosteal implants:

Subperiosteal implants were already introduced in the 1940s. Of all

currently used devices, it is the type of implant that has had the longest period

of clinical trial. These implants are not anchored inside the bone, such as End

osseous Implants, but are instead shaped to "ride on" the residual bony ridge

of either the upper or lower jaw. They are not considered to be Osseo

integrated implants. Subperiosteal Implants have been used in completely

edentulous upper and lower jaws (El-Askary, 2008).

It indicated in a severely resorbed, edentulous upper and lower jaw bone

(more than 5mm in height) (Park et al., 2008).

This implant is custom-made to the individual jaw.

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II.3.3 Tran osseous implants:

These implants are not in use that much any more, because they necessitate

an extra oral surgical approach to their placement, which again translates into

general anesthesia, hospitalization and higher cost, without providing higher

benefits to the patient. In any case, these implants are used in mandibles only

and are secured at the lower border of the chin via bone plates. These were

originally designed to have a secure implant system, even for much resorbed

lower jaws (more than 6mm height, 5mm buccolingually) (Cranin, 2001).

II.4 other types of implants:

II.4.1 Endodontically endosteal implant: implants which are placed

through endodically treated teeth into the bone to stabilize these teeth when

there are periodontal problem around them (Park et al., 2008).

II.4.2 Intramucosal inserts: button like non implanted device that used to

stabilize full and partial maxillary or mandibular removable denture (Cranin,

2001).

II.5 Types of implant's surgical procedures:

II.5.1 Two-stage surgery: the implant is surgically placed into the bone then

(after 3-6monthes healing periods) abutment and crown are connected to

implant.

Two-stage surgery is sometimes chosen when a concurrent bone graft is

placed or surgery on the mucosa may be required for esthetic reasons (Add et

al., 2008).

II.5.2 One-stage surgery: in which implant and abutment are placed at one

time (with out healing period) (Addy et al., 2008).

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II.6 Parts of Dental Implant (Implant System):

II.6.1 Implant or Fixture part:

A part of implant system that embeds into jawbone to provide anchorage

and support to other implant parts. It also allows bone tissue to grow around

the implant to reduce the bone loss occurs after natural teeth are lost.

Implants come in many different shapes (e.g. tapered), lengths and widths.

Materials Used in fixture are: Titanium, ceramic or zirconia (Young and

Sloan, 2001).

II.6.2 Abutment part:

It’s a part that provides support for the restoration (fixed or removable

partial denture). It is also the interface between the restoration and the

implant part. The abutment is eventually screwed to the implant using its

screw driver to guide it into position.

Material used: Titanium, ceramic or zirconia (Young and Sloan, 2001).

II.6.3 Restoration part (fixed or removable partial denture):

It is the part that looks like a tooth. It may be fixed or removable

restoration. Fixed restoration usually made of porcelain fused to a metal alloy

(PFM), but also could be a full-metal or full-porcelain crown. The crown is

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attached either to the abutment or directly to the implant, which can be

screwed or cemented onto the abutment (Young and Sloan, 2001).

II.7 Material used in implants

Requirements for implant materials:

The ideal properties of dental implant material for supporting restorations

may be classified into two groups, physical and mechanical properties and

biocompatibility (Autor, 2003).

Physical and Mechanical properties:

Implant materials must be having a high yield strength which describes the

ability of implant to bear loads without buckling (excessive permanent

deformations). The yield strength also determines their ability to prevent

failure due to distortion under occlusal forces (Lausmaa, 2009).

A high modulus of elasticity is needed for implant to distribute forces to

surrounding bone tissue (Autor, 2003).

Fracture toughness indicates the ability of implant to resist fracture in the

presence of flow or damage on the surface of a roughened or plasma sprayed

implant, so that implant materials might have high fracture toughness (Mante

and Mante, 2002).

Biocompatibility:

The implant materials should be biocompatible. High corrosion resistance of

implant materials promotes biocompatibility. Also they must have an osteo

inductive surface to accelerate Osseo integration (Cook et al., 2008).

Now day the most used materials in implant logy are:

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Titanium: the main material which consist the implant fixture and abutment.

Ceramics: used for all implant system (parts).

Zirconia: used for all implant system (parts) (Young and Sloan, 2001).

II.7.1 Titanium

Commercially pure titanium (CP-Ti) has been used since 1950 in

applications requiring high corrosion resistance, good shape-ability and good

welding capacity. CP-Ti is available in different grades, with different

amounts of impurities such as carbon, hydrogen, iron, nitrogen and oxygen.

Some CP-Ti alloys can incorporate small amounts of palladium (Ti-0.2Pd)

and nickel-molybdenum (Ti-0.3Mo-0.8Ni). These elements report

improvements to the mechanical resistance. Generally speaking, CP-Ti’s

main impurities consist of more than 1000 ppm of oxygen, iron, nitrogen,

carbon and silicon (Brunski, 2000).

The limitations related to monophasic-alpha- alloys, such as CP-Ti grades,

with low mechanical strength, low formability and fragility lead to the study

and development of biphasic-alpha/beta- alloys, such as Ti-6Al-4V. Ti-6Al-

4V is produced in a number of formulations. The oxygen content may vary

from 0.08 to more than 0.2 per cent, the nitrogen content may be adjusted up

to 0.05 per cent, the aluminum content may reach 6.75 per cent, and the

vanadium content may reach 4.5 per cent (Lausmaa, 2009).

Higher content of these elements, particularly oxygen and nitrogen lead to

the higher strength and the lower the ductility and fracture toughness

(Brunski, 2000).

Ti-6Al-4V is a useful material for surgical implants because of its low

modulus of elasticity, good tensile and fatigue properties, and biological

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compatibility. It is used for bone screws and for partial and total hip, knee,

elbow, jaw, finger, and shoulder replacement joints. It is not used as much as

CP-Ti in dental applications (implant) because loads borne by the dental

implants are not as high as in other surgical applications and Ti-6Al-4V has

less corrosion resistant than CP-Ti (Autor, 2003).

II.7.1.1 Classification of CP-Ti

American Standards for Testing and Materials (ASTM) recommends CP-Ti

alloys are classified in four different grades according to their mechanical

properties (Campbel et al., 2006):

II.7.1.1 a Grade-1 CP-titanium

The ASTM Grade-1 CP-titanium is chemically the purest. As a

consequence of its low content of interstitial elements, it has the lowest

mechanical strength and the highest ductility, shape-ability and workability,

at room temperature, of all four grades (Batzer et al., 2008).

Grade-1 is used when maximum workability is required and the main

concern is to increase corrosion resistance by reducing both the iron content

and the interstitial elements. It has an excellent behavior from high oxidizing

environments to medium reducing ones, including chlorides (Niinomi, 2008).

II.7.1.1b Grade-2 CP-titanium

The ASTM Grade-2 CP-titanium is an ideal material for industrial

applications because it has guaranteed yield strength of a minimum value of

275 MPa. This strength is comparable to the annealed austenitic stainless

steel and is used in applications where excellent ductility and shape-ability

are needed. Grade-2 has low contents of interstitial elements and as a

consequence the corrosion resistance is also improved. It also has good

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impact properties at low temperatures and an excellent wear and corrosion

resistance to saline solutions (Long and Rack, 2006).

II.7.1.1c Grade-3 CP-titanium

The ASTM Grade-3 CP-titanium has excellent corrosion resistance in

environments going from high oxidizing to medium reducing, including

chlorides. It has excellent specific strength and this is why Grade-3, as well as

other titanium alloys, is halfway between high resistance steels and light

aluminum alloys. It has good fracture toughness to impact at low

temperatures. The maximum limits in weight of iron in Grade-3 are lower

than in Grade-4 (0.3 per cent vs. 0.5 per cent) and Grade-3 has the second

highest value of oxygen (0.35 per cent) of the four grades. Only Grade-4 has

greater mechanical strength than Grade-3 (Bonollo and Gramgna, 2002).

II.7.1.1d Grade-4 CP-titanium: (most commonly used in dental implant)

The ASTM Grade-4 CP-titanium has the highest values of mechanical

strength of the four grades. It also has acceptable ductility and conformation.

The benefits of high mechanical strength and low density of Grade-4 can be

maintained up to moderate temperatures. Its specific mechanical strength is

superior to that of stainless steel AISI-301 even at temperatures above 315°C

(Long and Rack, 2006).

Grade-4 has excellent corrosion-fatigue resistance in saline solutions. The

stress required to attain fracture after a few million cycles is 50 per cent

higher than the stress needed for stainless steel AISI-341 (Autor, 2003).

Grade-4 has the highest content in weight of oxygen (0.40 per cent) and

iron (0.50 per cent) of the four CP-Ti grades (Cook et al., 2008).

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Grade-4 is available in all forms of production and can be mechanized,

molded, welded and cold-worked. All these processes can be performed at

room temperature but hot production (between 150 and 425°C) is normally

used to reduce the elastic recovery and energy required during production.

This method is used to produce complex shapes during manufacturing.

Grade-4 has an annealed equiaxed structure in all its forms of production (De

Groot et al., 2000).

II.7.1.2 improving the reliability of implants Osseo integration:

As has already been said, there are two technological approaches to

optimizing the fixation of dental implants: one changes the topography of the

implant surface and the other changes its chemical composition. The former

approach increases the surface roughness in order to improve long-term

fixation by assuring better implant-bone interlocking (Long and Rack, 2006).

As a consequence, the bone grows as near as possible to the titanium

implant and no fibrous tissue is observed between the implant and the bone

when using optical microscopy, as first given by Brånemark 1977 (Peltola et

al., 2009). In fact, this kind of interlocking should lead to a structural and

functional direct connection between the bone tissue and the surface of an

implant, i.e. Osseo integration of the dental implant (Boyer and Hall, 2003).

This includes several processes, among which is the shot-blasting treatment.

With this method a direct link between the material and the surrounding bone

is not produced. Nonetheless, in this case a thin coating of fibrous tissue,

which can only be observed by electron microscopy, surrounds the implant

(Lausmaa, 2009).

As to the latter approach, the absence of a chemical link between titanium

and bone tissue has lead to the development of several techniques to try to

modify the chemical composition of the implant. Some of these processes

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include electrophoresis deposition (Ducheyne et al., 2000), plasma-spray

(Loh, 2009), ion beam or radiofrequency attack (Hero et al., 2004), laser

ablation (Cleries, 2009) and isostatic pressing (Lausmaa, 2009). However,

none of these has been able to produce coatings chemically linked to the

substrate (Cook et al., 2008).

Nowadays, a common procedure used for clinical applications is the

coating of hydroxyapatite by plasma-spray (De Groot et al., 2000), since

hydroxyapatite is a bioactive material (Bruijn et al., 2001), a hydroxyapatite-

coated implant can stimulate bone cellular activity without any foreign-body

reaction, offering the possibility of complete Osseo integration of the implant.

However, one drawback to this method is the high temperature needed during

the plasma projection of the hydroxyapatite onto the titanium surface (De

Groot et al., 2000). Others are related to the difficult control of the chemical

composition, the crystalloid and the physical structure of the hydroxyapatite

during deposition because of its thermal instability (Bruijn et al., 2001).

II.7.2 Ceramics

Ceramic is a silicate in nature and may be defined as a combination of one

or more materials with a non metallic element, usually oxygen (Hiyasat et al.,

2009).

Dental ceramics were first used in dentistry in the late 1700s. Porcelain

jacket crowns were developed in the early 1900s. They consisted of

feldspathic or aluminous porcelain baked on a thin platinum foil and can be

considered the ancestors of all-ceramic crowns (Tinschert et al., 2010).

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II.7.2.1 Composition of Ceramics:

The quality of any ceramic depends on the choice of ingredients; correct

proportioning of each ingredient, and control of the firing procedure (Oden et

al., 2001).

Only the purest ingredients are used in the manufacture of dental porcelains

because of the stringent requirements of color, toughness without brittleness,

insolubility, and translucency, as well as the desirable characteristics of

strength and thermal expansion. In many instances, the manufacturer must

formulate a product that is a compromise (Odman and Anderson, 2001).

The average of dental porcelain wills there fore contains:

Feldspar (75-85%): the main component of feldspar is silicon oxide.

When undergo fusion, it form the glassy material which gives the porcelain

its translucency (Boening et al., 2000).

Quartz (silica) (12-22%): which contributes stability to the mass during

firing by provide frame work to other ingredients (Craig, 2008).

Kaolin (3-5%): it is clay which gives the porcelain its properties of

opaqueness and when mixed with water it becomes sticky material that binds

the other particles together when the porcelain unfired (Boening et al., 2000).

Coloring pigments: This is added to porcelain mixture in small quantities

to obtain delicate shades necessary to imitate the color of natural teeth (Craig,

2008).

II.7.2.2 Properties of dental porcelain (as general):

The properties of dental ceramic differ according to types of dental ceramic

but as a general they divided into:

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II.7.2.2a Physical properties: as (shrinkage, thermal properties).

II.7.2.2b Mechanical properties: (compressive strength, shear strength

and tensile strength).

II.7.2.2c Bio compatibility

II.7.2.2d Esthetical properties

II.7.2.2a Physical properties:

Shrinkage: linear shrinkage of glazed porcelain approximately 14% for low

fusion porcelain and 11.5% for high fusion porcelain (Chai et al., 2000).

Thermal properties: it has low thermal conductivity of 0.0030˚C/c and

12×10¯6/c coefficient of thermal expansion (Chai et al., 2000).

II.7.2.2b Mechanical properties:

Strength: it is deferent according to type of dental porcelain but as general

dental porcelain range from 172-450MPa compressive strength, 110-230MPa

shear strength and 34-70MPa tensile strength. So that dental porcelains are

brittle materials because of the strength of the silicon-oxygen bond and the

absence of grain boundaries, so that the glassy matrices of dental porcelain

have high intrinsic tensile strength (Deville et al., 2005).

So that, to decrease brittleness of dental ceramic: the design of ceramic

dental restorations should also avoid stress raisers in the ceramic. Thus any

sharp edges (as incisal edges, cusps or even sharp angels in implant

abutment) can cause stress concentration and act as stress raiser must be

rounded (Ban and Nawa, 2008).

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Also, minimize the number of firing cycles of porcelain to prevent the

mismatch between the veneer and the core in thermal expansion coefficients

that produce stresses during cooling that are sufficient to cause immediate or

delayed crack formation in the porcelain (Hiyasat et al., 2009).

The ion exchange is creates very large residual compressive stresses by

placing Sodium-containing glass article in a bath of molten potassium nitrate,

potassium ions in the bath exchange places with some of the sodium ions in

the surface of the glass article and remain in place after cooling. Since the

potassium ion is about 35% larger than the sodium ion, the squeezing of the

potassium ion into the place formerly occupied by the sodium ion creates

very large residual compressive stresses (Fernandez et al., 2007).

The thermal tempering creates residual surface compressive stresses by

rapidly cooling (quenching). This rapid cooling produces a skin of rigid glass

surrounding a soft (molten) core. As the molten core solidifies, it tends to

shrink, but the outer skin remains rigid. The pull of the solidifying molten

core, as it shrinks, creates residual tensile stresses in the core and residual

compressive stresses within the outer surface (Wen et al., 2000).

A further, yet fundamentally different, method of strengthening glasses

and ceramics is to reinforce them with a dispersed phase of a different

material that is capable of hindering a crack from propagating through the

material (Odman and Anderson, 2001).

II.7.2.2c Biocompatibility:

Ceramic is biocompatible some in vitro studies have been performed in

order to obtain information about cellular behavior towards ceramic. They

found that the ceramic is not cytotoxic. Ceramic doesn't induced bacterial

colonization on its surface (Tinschert et al., 2010).

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II.7.2.2d Esthetical properties:

Ceramic is esthetic material, it can be produced in thin layer (1-2mm) and

easily to shaped to produce highly esthetic prosthesis. It has a color similar to

natural tooth. Ceramic shade (color) can modify by increase/decrease pigment

content to match natural tooth color (Mante and Mante, 2002).

Factor affecting the color of ceramics:

The principal reason for the choice of porcelain as a restorative material is

its aesthetic quality in matching the adjacent tooth structure in translucence

and color (Fernandez et al., 2007).

Perfect color matching is extremely difficult, if not impossible. The

structure of the tooth influences its color. Dentin is more opaque than enamel

and reflects light. Enamel is a crystalline layer over the dentin and is

composed of tiny prisms or rods cemented together by an organic substance.

The indices of refraction of the rods and the cementing substance are

different. As a result, a light ray is scattered by reflection and refraction to

produce a translucent effect, and a sensation of depth as the scattered light ray

reaches the eye. As the light ray strikes the tooth surface, part of it is

reflected, and the remainder penetrates the enamel and is scattered. Any light

reaching the dentin is either absorbed or reflected to be again scattered within

the enamel. If dentin is not present, as in the tip of an incisor, some of the

light ray may be transmitted and absorbed in the oral cavity. As a result, this

area may appear to be more translucent than that toward the gingival area

(Mante and Mante, 2002).

Light rays can also be dispersed, giving a color or shade that varies in

different teeth. The dispersion can vary with the wavelength of the light.

Therefore the appearance of the teeth may vary according to whether they are

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viewed in direct light, this phenomenon is called metamerism (Esquivel and

Anusavice, 2000).

While dental porcelains are pigmented by the inclusion of oxides to

provide desired shades. So that, . it is impossible to imitate such an optical

system perfectly. The dentist and/or laboratory technician can, however,

reproduce the esthetic characteristics sufficiently so that the difference is

conspicuous only to the trained eye (Holand et al., 2006).

II.7.2.3 Classification of Dental Porcelain:

The ceramics can be classified according to firing temperature to:

High fusion (1270-1450˚C): is used for manufacturing of dental porcelain.

Medium fusion (1050-1200˚C).

Low fusion porcelain (850-1050˚C) (Craig, 2008).

The ceramics can be classified according to manufacturing methods into

(development of dental porcelain restoration):

Ceramic-metal restorations.

All-ceramic restorations: which divide into: sintering, heat-pressing, slip-

casting, and machining all-ceramic restorations (Shah et al., 2008)?

II.7.2.3a Ceramic-metal restorations:

Ceramic-metal restorations consist of a cast metallic framework (or core)

on which at least two layers of ceramic are baked. The first layer applied is

the opaque layer, consisting of ceramic rich in opacifying oxides. Its role is to

mask the darkness of the oxidized metal framework to achieve adequate

esthetics, it also provides ceramic-metal bond. The next step is the buildup of

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dentin and enamel (most translucent) ceramics to obtain an esthetic

appearance similar to that of a natural tooth. Opaque, dentin and enamel

ceramics are available in various shades (Denry et al., 2010).

The alloys used for casting the substructure are usually gold-based

containing tin and indium. Gold-palladium, silver-palladium, and nickel-

chromium alloys were initially developed as lower-cost alternatives.

However, the recent steep increase in the price of palladium has changed the

palladium-containing alloys into a higher-cost alternative (Craig, 2008).

It is essential that the coefficient of thermal expansion of the veneering

ceramic (8.6 × 10-6/°K) be slightly lower than that of the alloy to ensure that

the ceramic is in slight compression after cooling. This will establish a better

resistance to crack propagation of the ceramic-metal restoration (Hiyasat et

al., 2009).

II.7.2.3b All-ceramic restorations: Several processing techniques are

available for fabricating all-ceramic crowns: sintering, heat-pressing, slip-

casting, and machining (Hiyasat et al., 2009).

II.7.2.3b1 sintering all-ceramic restorations:

Two main types of all-ceramic materials are available for the sintering

technique: alumina-based ceramic and leucite-reinforced ceramic. (Denry et

al., 2010).

● Alumina-based ceramic:

Aluminous core ceramic contained 40% to 50% alumina by weight. Alumina

has a high modulus of elasticity (350 GPa), high fracture toughness (3.5 to 4

MPa m0, 5), flexural strengths of about 138 MPa and shear strengths of 145

MPa (Spear and Holloway, 2008).

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Its dispersion in a glassy matrix of similar thermal expansion coefficient

leads to a significant strengthening of the core. It has been proposed that the

excellent bond between the alumina and the glass phase is responsible for this

increase in strength compared with leucite-containing ceramics (Guazzato et

al., 2005).

● Leucite-reinforced feldspathic porcelain:

A feldspathic porcelain containing up to 45% by volume tetragonal

leucite is available for the fabricating all-ceramic sintered restorations.

Leucite acts as a reinforcing phase; the greater leucite content leads to higher

flexural strength (104 MPa) and compressive strength (Esquivel and

Anusavice, 2000).

The large amount of leucite in the material also contributes to a high

thermal contraction coefficient. In addition, the large thermal contraction

mismatch between leucite (22 to 25 x l0-6/º C) and the glassy matrix (8 x 10-

6/◦C) results in the development of tangential compressive stresses in the

glass around the leucite crystals upon cooling which act as crack deflectors

and contribute to increased resistance of the weaker glassy phase to crack

propagation (Hagg et al., 2004).

● Magnesia-based core porcelain:

A high-expansion magnesia core material has been developed that is

compatible with the same dentin porcelains used for ceramic-metal

restorations (Filser et al., 2003).

The flexural strength of unglazed magnesia core ceramic is twice as high

(131 MPa) as that of conventional feldspathic porcelain (70 MPa), with an

average coefficient of expansion of 14.5 x 10-6/◦C (Tinschert et al., 2010).

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II.7.2.3b2 Heat-Pressed all-ceramic restorations:

Heat-pressing classically helps avoid large pores and promotes a good

dispersion of the crystalline phase within the glassy matrix. The mechanical

properties of many ceramic systems are maximized with high density and

small crystal size (Albakry et al., 2004).

● Leucite-based:

Leucite-based ceramics are available for heat-pressing. Leucite (KA1Si2O6

or K2O. A12O3 . 4SiO2) and used as a reinforcing phase in amounts varying

from 35% to 55%. Ceramic ingots are pressed between 1150 and 1180°C

(under a pressure of 0.3 to 0.4 MPa) into the refractory mold made by the

lost-wax technique (Raigrodski, 2006).

To ensure compatibility with the thermal expansion coefficient of the

veneering porcelain, the thermal expansion coefficient of the material for the

veneering technique (14.9 x 10-6/°C) is lower than that of the material for the

staining technique (18 x 10-6/°C) (Coelho et al., 2009).

The flexure strength of these ceramics (120 MPa) is about double that of

conventional feldspathic porcelains. The main disadvantages are the initial

cost of the equipment and relatively low strength compared with other all

ceramic systems (Tholey et al., 2009).

● Lithium disilicate-based:

These materials contain lithium disilicate (Li, Si2O2) as a major crystalline

phase. They are heat-pressed in the 890 to 920°C temperature range, using

the same equipment as for the leucite-based ceramics. The heat pressed

restoration is later layered with glasses of matching thermal expansion. The

final microstructure consists of about 60% elongated lithium disilicate

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crystals (0.5 to 5 micrometers long) dispersed in a glassy matrix (Chevalier,

2009).

The main advantage of the lithium disilicate-containing ceramics is their

superior flexural strength (350 MPa) and fracture toughness (3.2 MPa. mO.5),

which extend their range of applications (Kelly and Denry, 2008).

II.7.2.3b3 Slip-cast all-ceramic materials

● Alumina-based:

An alumina-based slip is applied to a gypsum refractory die designed to

shrink during firing. The alumina content of the slip is more than 90%, with a

particle size between 0.5 and 3.5 ym. After firing for 4 hours at 1100 °C, the

porous alumina coping is shaped and infiltrated with a lanthanum-containing

glass during a second firing at 1150ºC for 4 hours (Hannink et al., 2000).

After removal of the excess glass, the restoration is veneered using

matched-expansion veneer porcelain. This processing technique is unique in

dentistry and leads to a high-strength material due to the presence of densely

packed alumina particles and the reduction of porosity. The flexural strength

of this slip-cast alumina material is around 450 MPa. Because of the high

strength of the core, short span anterior fixed partial dentures can be made

using this process (Filser, 2003).

● Spinel- and zirconia-based:

Two modified ceramic compositions for this technique have been recently

introduced. One contains a magnesium spinel (MgAl2O3) as the major

crystalline phase with traces of alpha-alumina, which improves the

translucency of the final restoration (Von et al., 2005).

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The second material contains tetragonal zirconium and alumina. The spinel-

based material has a lower modulus of rupture than the alumina based

material, whereas the zirconium-based material has a reported flexural

strength neighboring 600 MPa (Larson et al., 2006).

II.7.2.3b4 Machinable all-ceramic material:

One system uses CAD/CAM (computer assisted design/computer assisted

machining) technology to produce restorations in one office visit. After the

tooth is prepared, the preparation is optically scanned and the image is

computerized (Andreiotelli et al., 2009).

The restoration is designed with the aid of a computer. The restoration is

then machined from ceramic blocks by a computer-controlled milling

machine. The milling process takes only a few minutes. Although convenient,

the CAD/CAM system is very expensive and its marginal accuracy is poor,

with values of 100 to 150 pm. bonding of the restorations with resin cements

may help compensate for some of the problems of poor marginal fit (Lupu

and Giordano, 2007).

Another system for machining ceramics is to form inlays, on lays, and

veneers using copy milling. In this system, a hard resin pattern is made on a

traditional stone die. This handmade pattern is then copied and machined

from a ceramic block using a pantographic device similar in principle to those

used for duplicating house keys. Again, marginal accuracy is a concern and

there are high equipment costs (Sailer et al., 2007).

A more recent system involves an industrial CAD/CAM process to produce

crowns. The die is mechanically scanned by the technician and the data is

sent to a workstation where an enlarged die is milled using a computer-

controlled milling machine. This enlargement is necessary to compensate for

the sintering (Andreiotelli et al., 2009).

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II.7.3 Zirconium

Zirconium is a chemical element with the symbol Zr and atomic number 40.

Its atomic mass is 91.224. It is a lustrous, grey-white, strong transition

material that resembles titanium. Zirconium is used as an alloying agent for

its strong resistance to corrosion. It is never found as a native metal; it is

obtained mainly from the mineral zircon, which can be purified with chlorine.

Zirconium was first isolated in an impure form in 1824 by Jöns Jakob

Berzelius (Krebs and Robert, 2008). Zirconium forms both inorganic and

organ-metallic compounds such as zirconium dioxide and zircon-ocene

dichloride, respectively (Lide and David, 2008).

Naturally occurring zirconium is composed of five isotopes. Zr 90, Zr91, and

Zr92 are stable. Zr 94 has a half-life of 1.10×1017 years. Zr96 has a half-life of

2.4×1019 years, making it the longest-lived radioisotope of zirconium. Of

these natural isotopes, Zr90 is the most common, making up 51.45% of all

zirconium. Zr96 is the least common, comprising only 2.80% of zirconium

(Audi et al., 2003).

28 artificial isotopes of zirconium have been synthesized, ranging in atomic

mass from 78 to 110. Zr93 is the longest-lived artificial isotope, with a half-

life of 1.53×106 years. Zr110, the heaviest isotope of zirconium, is also the

shortest-lived, with an estimated half-life of only 30 milliseconds.

Radioactive isotopes at or above mass number 93 decay by β−, whereas those

at or below 89 decay by β+. The only exception is Zr88, which decays by ε

(Audi et al., 2003).

Zirconium is a lustrous, grayish-white, soft, ductile, and malleable which is

solid at room temperature, though it becomes hard and brittle at lower purities

(Emsley, 2001).

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II.7.4 Zirconium

Zirconia is a crystalline dioxide of zirconium. Its mechanical properties are

very similar to those of metals and its color is similar to tooth color. In 1975,

Garvie proposed a model to rationalize the good mechanical properties of

zirconia, by virtue of which it has been called ‘‘ceramic steel’’ (Addison et

al., 2003).

At ambient pressure, unalloyed zirconia can assume three crystallographic

forms depending on the temperature. At room temperature and upon heating

up to 1170ºC, the symmetry is monoclinic (P21/c). The structure is tetragonal

(P42/nmc) between 1170 and 2370ºC and cubic (Fm¯3m) above 2370 ºC and

up to the melting point (Bind et al., 2005). The transformation from the

tetragonal (t) phase to the monoclinic (m) phase upon cooling is accompanied

by a substantial increase in volume (4.5%), sufficient to lead to catastrophic

failure. This transformation is reversible and begins at 950◦C on cooling.

Alloying pure zirconia with stabilizing oxides such as CaO, MgO, Y2O3 or

CeO2 allows the retention of the tetragonal structure at room temperature and

therefore the control of the stress-induced t→m transformation, efficiently

arresting crack propagation and leading to high toughness (Munoz et al.,

2003).

II.7.4.1 Biocompatibility of zirconia:

The first proposal of the use of zirconium oxide for medical purposes was

made in 1969 and concerned orthopedic application. ZrO2 was proposed as a

new material for hip head replacement instead of titanium or alumina

prostheses (Heffernan et al., 2002).

They evaluated the reaction upon placing ZrO2 in a monkey femur and

reported that no adverse responses arose (Kosmac et al., 2007).

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Orthopedic research focused on the mechanical behavior of zirconia, on its

wear, and on its integration with bone and muscle. Moreover, these first

studies were largely carried out in vivo because in vitro technology was not

yet sufficiently advanced. Prior to 1990, many other studies were performed,

in which zirconia was tested on bone and muscle without any unfavorable

results (Ardlin, 2002).

Since 1990, in vitro studies have also been performed in order to obtain

information about cellular behavior towards zirconia. In vitro evaluation

confirmed that ZrO2 is not cytotoxic. Uncertain results were reported in

relation to zirconia powders that generated an adverse response. This was

probably due to zirconium hydroxide, which is no longer present after

sintering so that solid samples can always be regarded as safe (Heffernan et

al., 2002).

Mutagen city was evaluated by Silva and by Covacci, and both reported

that zirconia is not able to generate mutations of the cellular genome; in

particular, mutant fibroblasts found on ZrO2 were fewer than those obtained

with the lowest possible oncogenic dose compatible with survival of the cells

(Chevalier, 2006).

Moreover, zirconium oxide creates less flogistic reaction in tissue than other

restorative materials such as titanium. This result was also confirmed by a

study about peri-implant soft tissue around zirconia healing caps in

comparison with that around titanium ones. Inflammatory infiltrate, micro

vessel density, and vascular endothelial growth factor expression were found

to be higher around the titanium caps than around the ZrO2 ones (Kosmac et

al., 2007).

Also, the level of bacterial products, measured with nitric oxide synthase,

was higher on titanium than on zirconium oxide. Zirconia can up- or down-

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regulate expressions of some genes, so that zirconia can be regarded as a self-

regulatory material that can modify turnover of the extra cellular matrix

(Kosmac et al., 2007).

II.7.4.2 Mechanical properties of Zirconia (generally):

Zirconia has mechanical properties similar to those of stainless steel. Its

resistance to traction can be as high as 900-1200 MPa and its compression

resistance is about 2000 MPa (Oblak et al., 2004). Cyclical stresses are also

tolerated well by this material. Applying an intermittent force of 28 kN to

zirconia substrates, Cales found that some 50 billion cycles were necessary to

break the samples, but with a force in excess of 90 kN structural failures of

the samples occurred after just 15b cycles. Surface treatments can modify the

physical properties of zirconia. Exposure to wetness for an extended period of

time can have a detrimental effect on its properties. This phenomenon is

known as zirconia ageing. Moreover, also surface grinding can reduce

toughness. Kosmac confirmed this observation and reported a lower mean

strength and reliability of zirconium oxide after grinding (Dalskobler et al.,

2007).

II.7.4.3 Different types of zirconia ceramics available for dental

applications:

Although many types of zirconia-containing ceramic systems are currently

available, only three are used to date in dentistry. These are yttrium cation-

doped tetragonal zirconia polycrystals (3Y-TZP), magnesium cation-doped

partially stabilized zirconia (Mg-PSZ) and zirconia-toughened alumina (ZTA)

(Guazzato et al., 2005).

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II.7.4.3a 3Y-TZP

Biomedical grade zirconia usually contains 3mol% yttria (Y2O3) as a

stabilizer (3Y-TZP). While the stabilizing Y3+ cations and Zr4+ are randomly

distributed over the cationic sites, electrical neutrality is achieved by the

creation of oxygen vacancies (Oh and Anusavice, 2007).

3Y-TZP has been used to manufacture femoral heads in total hip

replacement prostheses since the late eighties but its use in orthopedic surgery

has since been reduced by more than 90%, mostly due to a series of failures

that occurred in 2001. 3Y-TZP is available in dentistry for fabrication of

dental crowns and fixed partial dentures and implant's abutment (Gamborena

and Blatz, 2006).

The restorations are processed either by soft machining of presintered

blanks followed by sintering at high temperature, or by hard machining of

fully sintered blocks (Patiket et al., 2004).

The mechanical properties of 3Y-TZP strongly depend on its grain size.

Above a critical grain size, 3Y-TZP is less stable and more susceptible to

spontaneous t→m transformation whereas smaller grain sizes (<1m) are

associated with a lower transformation rate .Moreover, below a certain grain

size (<0.2m), the transformation is not possible, leading to reduced fracture

toughness (Piwowrczyk et al., 2005). Consequently, the sintering conditions

have a strong impact on both stability and mechanical properties of the final

product as they dictate the grain size. Higher sintering temperatures and

longer sintering times lead to larger grain sizes (Coli and Karlsson, 2004).

Currently available 3Y-TZP for soft machining of dental restorations utilizes

final sintering temperatures varying between 1350-1550ºC depending on the

manufacturer. This fairly wide range of sintering temperatures is therefore

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likely to have an influence on the grain size and later the phase stability of

3Y-TZP for dental applications (Piwowrczyk et. al., 2005).

Chevalier demonstrated that the presence of cubic zirconia is not desirable

in 3Y-TZP for biomedical applications and is caused by uneven distribution

of the yttrium stabilizer ions. The cubic grains are enriched in yttrium while

the surrounding tetragonal grains are depleted and therefore less stable

(Chevalier et al., 2006).

As mentioned before, restorations produced by soft machining are sintered

at a later stage (i.e. following the forming steps), this process prevents the

stress-induced transformation from tetragonal to monoclinic and leads to a

final surface virtually free of monoclinic phase unless grinding adjustments

are needed or sandblasting is performed. Most manufacturers of 3Y-TZP

blanks for dental applications do not recommend grinding or sandblasting to

avoid both the t→m transformation and the formation of surface flaws that

could be detrimental to the long-term performance, despite the apparent

increase in strength due to the transformation-induced compressive stresses

(Larson et al., 2006).

In contrast, restorations produced by hard machining of fully sintered 3Y-

TZP blocks have been shown to contain a significant amount of monoclinic

zirconia. This is usually associated with surface micro cracking, higher

susceptibility to low temperature degradation and lower reliability (Mclarn

and Giordano, 2005).

The microstructure of 3Y-TZP ceramics for dental applications consists of

small equiaxed grains (0.2–0.5m in diameter) depending on the sintering

temperature. The mechanical properties are well above those of all other

available dental ceramics, with a flexural strength in the 800–1000MPa range

and fracture toughness in the 6–8MPam0.5 range. The Weibull modulus

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strongly depends on the type of surface finish and the processing conditions

(Gamborena and Blatz, 2006).

II.7.4.3b Glass-infiltrated zirconia-toughened alumina (ZTA):

Another approach to advantageously utilize the stress induced

transformation capability of zirconia is to combine it with an alumina matrix,

leading to a zirconia-toughened alumina (ZTA). One commercially available

dental product, In-Ceram® Zirconia® (VidentTM, Brea, CA), was developed

by adding 33 vol. % of 12mol% ceria stabilized zirconia (12Ce-TZP) to In-

Ceram® Alumina® (Raigrodski et. al., 2004).

In-Ceram® Zirconia® can be processed by either is slip casting or soft

machining. One of the advantages of the slip-cast technique is that there is

very limited shrinkage. However, the amount of porosity is greater than that

of sintered 3Y-TZP and comprises between 8 and 11% (Luthy et al., 2005).

This partially explains the generally lower mechanical properties of In-

Ceram® Zirconia® when compared to 3Y-TZP dental ceramics. It should be

pointed out, that Ce-TZP ceramics usually exhibit better thermal stability and

resistance to low temperature degradation than Y-TZP under similar thermo-

cycling or aging conditions (Stuart et al., 2007).

II.7.4.3c Magnesia partially stabilized zirconia (Mg-PSZ)

Although a considerable amount of research has been dedicated to magnesia

partially stabilized zirconia (Mg-PSZ) for possible biomedical applications,

this material has not been successful due mainly to the presence of porosity,

associated with a large grain size (30–60m) that can induce wear .The

microstructure consists of tetragonal precipitates within a cubic stabilized

zirconia matrix. The amount of MgO in the composition of commercial

materials usually ranges between 8 and 10mol% (Larson et al., 2006).

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Due to the difficulty of obtaining Mg-PSZ precursors free of SiO2,

magnesium silicates can form that lower the Mg content in the grains and

promote the t→m transformation. This can result in lower mechanical

properties and a less stable material. Denzir-M® (Dentronic AB) is an

example of Mg-PSZ ceramic currently available for hard machining of dental

restorations (Tinschert et al., 2010).

In esthetically demanding anterior regions, restoring a single-tooth space

with an implant-supported crown can be a challenge for clinicians (Zarb et

al., 2004).

Success of implant depends not only on a successful Osseo integration and

an implant's functional load-bearing capacity, but also on the harmonious

integration of a crown into the dental arch. For highly esthetic anterior

locations in the dental arch, especially in the patients with a high lip line,

implant-supported single-tooth restorations are subjected to the most exacting

requirements, including optimal implant and superstructure positioning.

(Tischler, 2004).

Dental implants and abutments are usually fabricated from commercially

pure titanium because of its well-documented biocompatibility and

mechanical properties (Adell et al., 1999).

Despite the numerous improvements in the fabrication and design of metal

abutments, still there's remains a risk of the metal components being visible

when such abutments are used. Even when placed subgingivally, a dull gray

background may give the soft tissue an unnatural bluish appearance

especially under all ceramic crowns. The presence of a gray gingival

discoloration may be attributed to a thin gingival tissue thickness in the area

around the abutment that is incapable of blocking reflective light from the

metal abutment surface (Glauser et al., 2004).

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Hence, for achieving optimal mucogingival esthetics; ceramic&/or

zirconium abutments were developed (Yildirim et al., 2000). Currently,

ceramic abutments are fabricated out of two high-strength ceramic materials:

a densely sintered high-purity alumina ceramic (Al2O3) and a zirconium

ceramic (ZrO2). Both materials have improved optical and mechanical

properties and demonstrate differences in their microstructure and mechanism

against flaw propagation (Wael et al., 2006).

II.8 Development of ceramic abutments

The first ceramic abutment -Ceramic Core-was introduced in 1993 in small

and large diameters (not commercially available). The abutment was a

prototype of alumina ceramic with resistance to shearing forces that reached

values up to those of the metal–ceramic crowns (Suzuki, 2008). Compared to

metal abutments, these new abutments offered optically favorable

characteristics, low corrosion potential, high biocompatibility, and low

thermal conductivity (Wong et al., 2004).

On the other hand, restorations made out of such ceramic cores were

weaker when compared to metal–ceramic restorations. Such controversies led

to further investigations into new designs and materials for ceramic

abutments. Custom-made ceramic abutments were fabricated using alumina

blocks and milled on a coping milling machine. The abutments showed

improved values for resistance to fracture but they were still weaker than the

CeraOne- abutments (Killer et al., 2004).

Another step toward perfecting the overall esthetic outcome was taken with

the development of the customizable CerAdapt- abutment. The abutment was

made of pure, highly sintered aluminum oxide and demonstrated significantly

improved resistance compared to previous abutments. It was indicated for the

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fabrication of implant-supported single crowns and short-span fixed partial

dentures in both anterior and premolar regions (Wennerberg et al., 2007).

II.9 Contemporary Ceramic abutment:

Today, the majority of implant manufacturers offer ceramic abutments. The

abutments are available in pre-fabricated or customizable forms and can be

prepared in the dental laboratory either by the technician or by utilizing

computer-aided design ⁄ computer-aided manufacturing techniques (Kim et

al., 2007).

The materials of preference are densely sintered high-purity alumina

(Al2O3) ceramic and yttria (Y2O3) -stabilized tetragonal zirconia poly-crystal

ceramics (Stuart et al., 2007).

These high-strength ceramics have improved mechanical properties.

Alumina ceramic has a flexural strength of 400 MPa, a fracture toughness

value between 5 and 6 MPa ⁄ m0.5, and a modulus of elasticity of 350 GPa

(Luthy et al., 2005). The yttria stabilized zirconia ceramic has twice the

flexural strength of alumina ceramic (900–1400 MPa), a fracture toughness of

up to 10 MPa ⁄ m0.5, and a modulus of elasticity value of 210 GPa (Strub

and Gerds, 2003).

Compared to alumina ceramic, the enhanced strength of zirconia (ZrO2) can

be explained by micro structural differences, such as higher density, smaller

particle size, and polymorphic mechanism against flaw propagation (Scherrer

et al., 2001). The main reason for the superior resistance of zirconia lies in

the stabilizing effect of yttria, which allows the processing of zirconia in the

metastable tetragonal crystalline structure at room temperature (18°C–23°C).

The tetragonal phase at room temperature allows for transformation to the

monoclinic phase under stress and represents an efficient mechanism against

flaw propagation (Wennerberg et al., 2007). The transformation results in a

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compressive stress as the result of volume expansion and slows down further

crack propagation, resulting in improvement of the mechanical properties (i.e.

transformation toughening) (Yildirim et al., 2003). Alumina abutments are

composed of 99.5% pure alumina ceramic. These abutments provide certain

aesthetic advantages when compared to the more whitish zirconia abutments.

In addition, the alumina ceramic is easier to prepare; this saves time during

definitive preparation, which is usually performed intra orally. The problems

presented by alumina abutments include their radio opalescence at the time of

radiographic examination and their weak resistance to fracture. In this

context, it is commonly agreed that ceramic abutments should show proper

resistance against the masticators forces raised during chewing or swallowing

(Kohal et al., 2008).

From the previous we find that the titanium has good mechanical properties

and biocompatible for dental implant (fixture and abutment), but has some

esthetic problem (when uses as abutment) especially in anterior region, so

that ceramic abutments use to over come this problem. Before performing in

vivo studies or applying these materials for clinical use, in vitro tests should

be undertaken to prove materials' applicability and performance. These tests

can be performed in a short period of time and have the advantages of

reproducibility and the possibility of standardizing test parameters (Wael et

al., 2006).

In May 1995 Tripodakis et al. were compared the strength and mode of

failure of three different designs of custom-made all-ceramic implant

abutments fabricated by milling of In-Ceram sintered ceramic blocks with the

conventional CeraOne system under static load, and found that there is no

significant difference and the weakest link in the all-ceramic single implant

restorations was the abutment screw in which the bending began at

approximately 190 N.

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In 2002 Hye-Won Cho et al. were compared five different abutment-crown

combinations for single implant-supported restorations regarding their

capabilities to withstand loads, and concluded that all-ceramic crowns on the

milled ceramic abutments were weaker than the metal-ceramic crowns on the

titanium abutments under oblique loading.

In September 2003 Paulino Castellon et al. were compared (in vitro)

fracture resistance of all ceramic system (zirconia abutment covered with

alumina crown) and metal implant system (titanium abutment with porcelain

fused to metal crown) in anterior region, and reached to the metal system has

higher fracture resistance than ceramic one but ceramic system with in

clinical acceptance.

In October 2003 Murat Yildirim was quantified the fracture load of Al2O3

and ZrO2 abutments restored with glass-ceramic crowns for anterior region,

and concluded that both all-ceramic abutments exceeded the established

values for maximum incisal forces reported in the literature (90 to 370 N).

The ZrO2 abutments were more than twice as resistant to fracture as the

Al2O3-abutments.

In December 2003 Henriksson et al. were evaluated the clinical

performance of customized ceramic single-implant abutments in combination

with two different techniques for fabricating crowns, and found the all

implants and restorations were still in function after 1 year. So that the short-

term results indicate that customized ceramic abutments are successful and

have comparable function, regardless of fabrication method.

In March 2004 Philip Leong Biow was compared the esthetic outcome of

replacing the same left maxillary central incisor with 2 types of implant-

supported restorations, a zirconia abutment with a Procera crown and a

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custom metal abutment with metal ceramic crown. There were only subtle

differences noted and both restorations yielded a satisfactory result.

In May 2004 Glauser et al. were evaluated clinically an experimental

implant abutment made of densely sintered zirconia with respect to peri-

implant hard and soft tissue reaction as well as fracture resistance over time.

They found that Zirconia abutments offered sufficient stability to support

implant-supported single-tooth reconstructions in anterior and premolar

regions. The soft and hard tissue reaction toward zirconia was favorable.

In 2005 F. Butz was compared titanium-reinforced ZrO2 and pure Al2O3

abutments regarding their outcome after chewing simulation and static

loading on central incisor (in vitro study), and found that titanium-reinforced

ZrO2 abutments perform similar to metal abutments, and can therefore be

recommended as an aesthetic alternative for the restoration of single implants

in the anterior region. All-ceramic abutments made of Al2O3 possess less

favorable properties.

In Jan 2006 Peter Gehrke et al. were determine the fracture strength of

zirconium implant abutments and the torque required to unfasten the retaining

screw before and after applying cyclic loading to the implant-abutment

assembly. The dynamic behavior and stress distribution pattern of zirconium

abutments were also evaluated, and they concluded that zirconium implant

abutments exceeded the established values for maximum incisal bite forces

reported in the literature, and tightly fit into the titanium implant after several

millions of loading cycles.

In February 2006 Wael Att was evaluated fracture resistance of zirconia,

alumina oxide and titanium abutments covered with alumina all-ceramic

restorations in vitro study in anterior region, and he found that all 3 implant-

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supported restorations have the potential to withstand physiologic incisal

forces applied in the anterior region

In May 2006 Wael Att was evaluated the fracture resistance of single-tooth

implant-supported all-ceramic restorations, composed of zirconium dioxide

all ceramic restorations on different implant abutments, and to identify the

weakest component of the restorative system, and he found that all tested

implant-supported restorations have the potential to withstand physiological

occlusal forces applied in the anterior region. Because of the low fracture

resistance values of group Al, the combination of zirconia crowns and

alumina abutments should carefully be considered before clinical application.

In May 2007 Anders Sundh were evaluate the bending resistance of

implant-supported CAD/CAM-processed restorations made out of zirconia or

manually shaped made out of reinforced alumina, and they concluded that

the all ceramic abutments and copies exhibited values that were equal or

superior to that of the control and exceeded the reported value, up to 300 N,

for maximum incisal bite forces.

In November 2008 Aramouni et al. were evaluated the fracture resistance

and failure location of single-tooth, implant-supported, all-ceramic

restorations on different implant abutments (ZrO2, Ti and ceramic abutment)

subjected to a maximum load, and found that The zirconium oxide (ZrO2)

ZiReal and titanium (UCLA) abutments on the 3i Certain implants had

statistically significantly higher fracture loads than those recorded for the 3i

Ceramic Blank abutments on the SLA ITI Straumann implant.

In Jan 2009 Adatia et al. were assessed the effect of different degrees of

clinical reduction of zirconia abutments on the failure load of clinical

assemblies (in vitro study), and found that all fractures occurred at the

interface where the abutment was connected to the analog. The preparation of

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zirconia abutments did not significantly impair the fracture resistance of

simulated implant assemblies. All implant abutments fractured at rates higher

than the maximum incisal forces (90-370 N) estimated to occur in the anterior

region of the mouth.

In September 2009 Sailer , made manual searching to identify randomized-

controlled clinical trials and prospective and retrospective studies providing

information on ceramic and metal abutments with a mean follow-up time of

at least 3 years. Patients had to have been examined clinically at the follow-

up visit. Assessment of the identified studies and data abstraction was

performed independently by three reviewers. He reached to the all-ceramic

crowns supported by ceramic abutments exhibited similar annual fracture

rates as metal-ceramic crowns supported by metal abutments.

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Aim of the Work

The aim of this study is to compare Zirconia & Alumina abutments versus

conventional Titanium abutments supporting all ceramic crowns in vitro by:

1) Measuring fracture resistance of crown and abutment.

2) Analyzing mode of failure using SEM.

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IV.1 Materials:

Table IV.1: Detail description of materials utilized in this study:

Material Composition Manufacturer Badge No. (if

present) Analogue (laboratory implant)

Commercially pure titanium (CP-Ti) grade 4

Biohorizons, USA

TM0900471

Ti abutments Commercially pure titanium (CP-Ti) grade 4

Biohorizons, USA

TM201402

ZrO2 abutments Y2O3_partialy stabilized ZrO2

Biohorizons, USA

TMR381565

Al2O3 abutments Densely sintered highly purity Al2O3

Biohorizons, USA

TMR38112

All ceramic crowns, (IPS e-max )

Leucite reinforced heat pressed glass –ceramic (IPS Empress staining technique)

Ivolar-Vivadent, Schann, Lietchtenstein

J25824

Adhesive cementation (Rely X ARC)

Universal self adhesive Resin Cement

Rely X, 3M ESPE, Seefeld, Germany

N126568

Hydrofluoric etching (ceramic etching gel 4.5 weight-% HF)

Ivolar-Vivadent, Schann.

E33956

Silanizations of pretreated ceramic surface (Monobond-S)

RelyXTM Ceramic Primer, 3M ESPE.

S322746

Tribochemically silicoated with a modified Rocatec-method 110-μm grain size Rocatec Plus

3M ESPE, Seefeld, Germany

___________

Rubber base impression material (Oramadent)

Base: Polysulfide polymer Titanium dioxide, zinc sulfate.

Oramadent, Morcalieri, Italy

2310024

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Accelerator: Lead dioxide Dibutyl

Stone Calcium sulfate hemi hydrate.

SO68731

Die spacer (Pico Fit)

Acetone, Butyl acetate, Methyl Ethyl Ketone and mineral spirits

Pico Fit, Pearson lab, Sylmar California, USA

T266629

Investment material (Speed investment)

Silicon dioxide, a-calcium sulfate hemi hydrate and sodium chloride

Speed investment, Ivolar-Vivadent, Schann, Lietchtenstein

OT102209

Wax natural and synthetic waxes, gums, fats, fatty acids, oils, natural and synthetic resins, and pigments

Yamahachi, Japan

1702051

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IV.2 Methods:

Forty eight analogues (resembling implant) with a diameter of 4 mm and

length of 13 mm1 (figure IV.1) represent missed root of upper first premolar

were used in this study. The implants were divided, according to the type of

abutments used into three groups (of 16 specimens each):

Group I: titanium abutments2 were used.

Group II: ready made zirconia abutments (ZrO2)3 were used.

Group III: ready made alumina abutments (Al2O3)3 were used.

All abutments (Ti, ZrO2 & Al2O3) had standard measurements: a deep

chamfer finish line of 1 mm depth and a total height of 6 mm &4mm

diameter (figure IV.II).

Figure (V.I): Implant analogue.

1 Biohorizons, USA. 2 Esthetic Titanium abutment, Biohorizons, USA. 3 Esthetic Ceramic abutment, Biohorizons, USA.

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Figure (IV.1): Titanium, zirconia and alumina abutment respectively.

IV.2.1 Crown fabrication:

Glass infiltrated ceramic crown system was used (IPS e-max press MO)4.

Each abutment was fixed in its analogues through titanium screw by screw

driver (figure IV.2). Then specimen (abutment and analogue) (figure IV.3)

was fixed in the model (represent missed premolar) (figure IV.4) and

impression was taken for full arch by rubber base impression material

(Oramadent)5 and casted in hard stone (Sheerer stone)6, then stone die was

prepared and sealer was applied to harden the surface and to protect the stone

die. However, the sealer layer must not lead to changes of the dimensions of

the stone die (figure IV.5).

4 IPS Empress staining technique, Ivolar-Vivadent, Schann, Lietchtenstein. 5 Oramadent, Morcalieri, Italy 6 Sheerer Stone: Ivolar-Vivadent, Schann, Lietchtenstein

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Die-spacer (Pico fit)7 was applied in two layers up to maximum 1 mm

from the preparation margin (spacer application 9-11 pm per layer) (we

would sure to consider the expansion of the investment material when

applying the spacer).

The wax-up was designed fully anatomical using an ash-free wax8 to

approximately 0.7 mm thickness, 6mm length and 4mm diameter (figure

IV.6).

The sprues (with 3-8 mm in length and 2.5-3 mm in width) were placed in

the direction of flow of the ceramic (45-60° to the investment ring base and

axial to wax pattern) and at the thickest part of the wax pattern in order to

achieve unimpeded flow of the viscous ceramic material.

\

Figure IV.2: Abutment fixed in its analogues through titanium screw by

screw driver.

7 Pico Fit, Pearson lab, Sylmar California, USA. 8 Yamahachi, Japan

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Figure IV.3: Titanium, Zirconia and Alumina abutment after fixation of their

analogue.

Figure IV.4: Upper model with abutment fixed in position of upper premolar.

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Figure IV.5: Stone die of abutment.

Figure IV.6: Wax pattern of glass ceramic crown.

Investing was carried out with IPS PressVEST Speed9 (figure IV.7). The

100 g investment ring base was selected and the corresponding IPS silicone

ring with matching ring gauge was used for investing purpose. The IPS

Silicone Ring was carefully positioned on the investment ring base without

damaging the wax objects. The IPS Silicone Ring was sitted flush on the

investment ring base,then investment material was mixed with a suitable

instrument for the fine investment of the cavity. Subsequently, the

investment ring was carefully filled with investment material up to the

9 Speed investment, Ivolar-Vivadent, Schann, Lietchtenstein.

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marking and positioned the ring gauge with a hinged movement (figure

IV.8). Finally allow the investment material to set without manipulating the

investment ring.

After the setting time of the investment material (30-45 min) finished, the

ring gauge and ring base were removed with a turning movement and

investment ring was carefully push out of the IPS Silicone Ring, then rough

spots on the bottom surface of the investment ring was removed with a

plaster knife.

Finally, investment block (with wax pattern) was preheated by tipping the

block with opining facing down towards the rear wall in the furnace at 950°C

temperature for 30 min (preheating was started from room temperature and

rise to750°C in which it still for 10 min then rised to 950°C) (figure IV.9).

Figure IV.7: IPS PressVEST Speed.

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Figure IV.8: Investment ring filled up to the marking and positioned the ring

gauge with a hinged movement.

Figure IV.9: Investment ring placed towards rare wall of preheating furnace.

Before the preheating cycle for the investment ring was ended, the

following preparations for pressing must be carried out:

A cold IPS e.maxAlox Plunger and a cold IPS e.max Press Ingot were

provided in the desired shade (figure IV.10). After that, the cold IPS e.max

Alox Plunger was dipped into the opening of the IPS e.maxAlox Plunger

Separator and keep ready for use, then the press furnace was turned on in

time. Finally the press program for IPS e.max Press was selected.

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Once the preheating cycle has been completed (no more than 1 minute for

these steps to prevent the investment ring from cooling down too much), the

investment ring was removed from the preheating furnace and the cold IPS

e.max Press ingot was inserted at the rounded, non-imprinted side into the hot

investment ring (the imprinted side should face upward to double-check the

ingot shade) (figure IV.11). The powder-coated cold IPS e.maxAlox plunger

was placed into the hot investment ring (figure IV.12). Then the completed

investment ring was placed at the center of the hot press furnace using the

investment tongs ring (figure IV.13). Finally, the selected press program

was started (at 950°C, 6 bar for 35min).

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Figure IV.10: A cold isolated IPS e-max Alox Plunger and a cold IPS e-max

Press ingot selected in the desired shades.

Figure IV.11: Insertion of cold IPS e-max Press ingot into hot investment

ring.

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Figure IV.12: Insertion of powder-coated IPS e-max Alox plunger into hot

investment ring.

Figure IV.13: Insertion of the hot, completed investment ring at the center of

the hot press furnace using investment tongs.

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After cooling to room temperature (approximately 60 minutes), the

investment ring may show cracks. These cracks developed (immediately

around the Alox plunger) during cooling as a result of the different

coeffecient of thermal expansions of the various materials (Alox Plunger,

investment material, and pressed materials).

the investment ring as follows: the length of the Alox plunger was marked

on the cooled investment ring (figure IV.14). Then the investment ring was

separated using a separating disk (figure IV.15).

Rough divestment was carried out with glass polishing beads at 4 bar (60

psi) pressure (figure IV.16). Then, fine divestment was carried out with

glass polishing beads at 2 bar (30 psi) pressure. Finally, ceramic residue on

the Alox plunger was removed with Al2O3 (type 100 microns).

After fine divestment, the reaction layer formed during the press procedure

was removed using IPS e.max press invex liquid (figure IV.17). (containing

>1% hydrofluric acid) by: the invex liquid was poured into a plastic cup.

Then, the pressed object was completely immersed into the invex liquid and

cleaned in an ultrasonic cleaner for at least 10 minutes and a maximum of 30

minutes. Subsequently, the object was cleaned under running water and blew

dry. Finally, the white reaction layer was removed using Al2O3 type 100 at

1-2 bar (15-30 psi) pressure (reaction layer might be completely removed to

avoid bubbles may develop, which may lead to bonding problems or even

cracks in the layering ceramic).

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Figure IV.14: The length of Alox plunger marked.

Figure IV.15: Separation of investment ring using separating disk at

predetermined point.

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Figure IV.16: Rough investment with glass polishing beads.

Figure IV.17: Removal of reaction layer.

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The area to be ground was wetted and used a fine diamond disk to cut the

sprues (overheating of the ceramic material might be avoid, so that low speed

and light pressure was recommended) (figure IV.18). Then, the attachment

points of the sprues were smoothed out. The spacer was removed prior to

placing the pressed crown on the die. The crown was placed on the die and

carefully adjust. The crown was blasted with Al2O3 at 1 bar (15 psi) and

cleaned under running water or with steam before applying the veneering

material.

A honey-combed firing tray and the corresponding support pins were used

to fire the restorations (figure IV.19). The top edges of the support pin were

covered with platinum foil or a small amount of IPS Object Fix Putty or flow

to prevent the crown from sticking to the pin. Any contamination on the

crown after cleaning was removed. After that, the wash firing was conducted

using IPS e.max Ceram Shades and Essence. The paste or powder was mixed

with the IPS e.max Ceram Stain Liquids (allround or longlife) to the desired

consistency. The wash was apllied in a thin coat on the entire crown. Finally,

the crown was entered furnace (EP 5000) at 770°C for 18 min (temperature

rised from room temperature to 403°C then still for 4min after that rised to

770°C).

Finally, the crown was glazed as following: IPS e.max Ceram Glaze was

extruded from the syringe and the material was thined to the desired

consistency using IPS e.max Ceram Glaze and Stain Liquid. The even layer

glazing material was applied on the entire outer surfaces of the crown.

fluorescent glaze (paste or powder) was applied to cervical areas. Finally,

the crown was entered furnace (EP 5000) at 750°C for 12min.

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Figure IV.18: The sprue cutted with fine diamond disk.

Figure IV.19: Hony combed firing tray.

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IV.2.2 Fixation of Specimen:

Afterwards, specimens were removed from the model then embedded in

special specimen holders using an auto polymerizing acrylic resin vertically

to the horizontal plane to simulate clinical conditions (figure IV.20). The

resin has a modulus of elasticity of approximately 12 GPa, which

approximates that of human bone (18 GPa) (Wael et. al., 2006).

IV.2.3 Cementation of Crowns:

To ensure maximum bond strength between crowns and the Al2O3, ZrO2

and Ti abutments a universal self adhesive resin (Rely-X ARC)10 was used

(figure IV-21). The internal surface of the glass-ceramic crowns was etched

with 4.5% hydrofluoric acid11 for 60 seconds (figure IV.22), carefully

cleaned with water spray application, and then dried by air for 30 seconds.

Al2O3, ZrO2 and Ti abutments were sandblasted12 for 60 seconds (figure

IV.23) (Blixt et al., 2001).

Resin was mixed according to the manufacturer's guidelines to lute the

crown to the abutment. Excess luting agent was removed with resin pellets.

Subsequently, a glycerin gel was applied to the abutment-crown interface to

prevent the formation of an oxygen-inhibited layer. Polymerization was

achieved with a high-performance polymerization light applied for 90

seconds on each surface (Wael et. al., 2006).

10 Relay X ARC, 3M ESPE, Seefeld, Germany 11 Ivolar-Vivadent, Schann. 12 Rocatec-method 110-μm grain size Rocatec Plus, 3M ESPE, Seefeld, Germany.

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Figure IV.20: Fixation of specimen in acrylic holder.

Figure IV.21: Rely-X adhesive cement.

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Figure IV.22: The internal surface of ceramic crown etched with Hf acid.

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Figure IV.23: Sandblasting of abutments.

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IV.2.4 Distilled water storage:

Each specimen was stored in bottle filled with distilled water for 24 hours

before application of thermal and mechanical loading (figure IV.24).

IV.2.5 Application of thermal and dynamic load:

After storage of distilled water, thermo cycling was done in thermo cycling

apparatus for 600 cycles from 5°C to 55°C with 2 min. dwell time, 10 sec.

transfer time (figure IV.25).Then the crown of specimen was covered with

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load stamp and subjected to maximum vertical load of 10 kg. with cyclic

frequency of 1.7 Hz for 240.000 cycles (figure IV.26) which correspond to

12 months of clinical service ( Kerjci et al, 1993).

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Figure IV.24: Specimen storage in distilled water.

Figure IV.25: Thermal cycling.

Figure IV.26: Dynamic Loading.

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IV.2.6 Measurement of fracture load:

All samples were individually & vertically mounted on a computer

controlled materials testing machine (Model LRX-plus; Lloyd Instruments

Ltd., Fareham, UK) (figure IV.28) with a load cell of 5 kN and data were

recorded using computer software (Nexygen-MT; Lloyd Instruments). The

samples were secured to the lower fixed compartment of the machine by

tightening screws. Load was applied with a custom made load applicator (A

steel rod with round tip 3 .6mm diameter) attached to the upper movable

compartment of the machine to contact the inclined planes of cusps. Tin foil

sheet was placed between the loading tip and the occlusal surface of crown

samples to achieve an even stress distribution and minimization of the

transmission of local force peaks (figure IV.27). Samples were statically

compression loaded until fracture at a crosshead speed of 1 mm/min. Failure

manifested by first crack sound and confirmed by sudden drop along the

load-deflection curves which were recorded with computer software

(Nexygen; Lloyd InstrumentsLtd).

IV.2.7 SEM analysis:

Specimens were scanned by electron microscope to evaluate the failure

mode. Specimens were coated with gold coating (SPI-Modules Vac/Sputter

Coater) which made as conductor for electron beam. Then, specimens were

scanned by electron microscope (JEOL-JSM-5200LV Japan) at 35 times

magnifications, then at 500 times magnifications.

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Figure IV.27: measurement of fracture resistance under testing machine

(Model LRX-plus; Lloyd Instruments Ltd).

Figure (IV.28): Universal Testing Machine (Lloyd Instruments Ltd).

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All test specimens survived 240.000 cycles in Dynamic Loading Machine

and 600 thermal cycles. No screw loosening was recorded. The lowest

fracture resistance value after oral simulation and application of the load to

fracture testing was observed in Al2O3 group 492.1N, whereas the highest

value was observed in Ti group 1007N.

Data analysis was performed in several steps. Initially, descriptive

statistics of fracture resistance test results for all groups including minimum,

maximum, mean, standard deviation, standard error and median. Comparison

of fracture resistance results for all groups was done by ANOVA test

followed by Newman Keuls post-hoc tests. Student t-test was done to detect

significance between paired groups. Statistical analysis was performed using

Graphpad Prism-4 statistics software for Windows. P values ≤ 0.05 are

considered to be statistically significant in all tests.

Descriptive statistics of fracture resistance test results for all groups are

presented in table (V.1) and graphically drawn in figure (V.1):

Table (V.1) Descriptive statistics of fracture resistance results for all groups

Ti ZrO2 Al2O3

Maximum 1007 812.1 676.9

Minimum 699.6 703.8 492.1

Mean 787.5 779 562.9

Std. Deviation 82.97 39.7 65.49

Std. Error 20.74 9.924 16.37

Median 781.5 800.8 548.2

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Figure (V.1): A column chart of fracture resistance mean values for all groups

It was found that Ti group recorded the highest fracture load mean value

(787.5 ± 82.97 N) followed by ZrO2 group (779.0 ± 39.7 N) while Al2O3

group recorded the lowest fracture load mean value (562.9 ± 65.49 N)

The difference between fracture resistance mean values for all groups was

statistically significant as revealed by ANOVA followed by Newman-keuls

post-hoc tests (p<0.05)

Table (V.2): One way ANOVA test comparing between fracture resistance

results for all groups

Group Mean ± SD ANOVA

Ti 787.5 ± 82.97 F P value

ZrO2 779.0 ± 39.7 61.02 <0.0001*

Al2O3 562.9 ± 65.49

*; significant (p<0.05)

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Ti vs. ZrO2

It was found that Ti group recorded higher fracture load mean value (787.5

± 82.97 N) than ZrO2 group (779.0 ± 39.7 N)

The t-test analysis showed non-significant difference between Ti and ZrO2

groups (t= 0.371; P > 0.05).

Table (V.3) Comparison of fracture resistance results (Mean±SD) between Ti

and ZrO2 groups

Mean±SD Difference t-test

Ti 787.5 ± 82.97 8.53 t-test P value

ZrO2 779.0 ± 39.7 0.371 0.7133 ns

*; significant (p<0.05) ns; significant (p>0.05)

Figure (V.2): A column chart of fracture resistance mean values for Ti and

ZrO2 groups

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Ti vs. Al2O3

It was found that Ti group recorded higher fracture load mean value (787.5

± 82.97 N) than Al2O3 group (562.9 ± 65.49 N)

The t-test analysis showed significant difference between both Ti and Al2O3

groups (t= 8.5; P < 0.05).

Table (V.4): Comparison of fracture resistance results (Mean±SD) between

Ti and Al2O3 groups

Mean±SD Difference t-test

Ti 787.5 ± 82.97 224.6 t-test P value

Al2O3 562.9 ± 65.49 8.5 <0.0001*

*; significant (p<0.05) ns; significant (p>0.05)

Figure (V.3): A column chart of fracture resistance mean values for Ti and

Al2O3 groups

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ZrO2 vs. Al2O3

It was found that ZrO2 group recorded higher fracture load mean value

(779.0 ± 39.7 N) than Al2O3 group (562.9 ± 65.49 N)

The t-test analysis showed significant difference between both ZrO2 and

Al2O3 groups (t= 11.26; P < 0.05).

Table (V.5): Comparison of fracture resistance results (Mean±SD) between

ZrO2 and Al2O3 groups

Mean±SD Difference t-test

ZrO2 779.0 ± 39.7 216.1 t-test P value

Al2O3 562.9 ± 65.49 11.26 <0.0001*

*; significant (p<0.05) ns; significant (p>0.05)

Figure (V.4): A column chart of fracture resistance means values for Al2O3

and ZrO2groups

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The location and mode of failure of the 3 test groups after exposure to

clinical simulation are reported in table (V.6):

Thermal and Dynamic

Loading

Static loading

Group Survival Failure Deflection

of crown

only

Abutment

screw

fracture

Abutme

nt &

crown

fracture

Neck

distorti

on

ZrO2 16 0 12 0 4 0

Al2O3 16 0 0 0 16 0

Ti 16 0 16 0 0 0

Table (V.6): Showed all test specimens were survived after aging, no screw

fracture (loosening) or neck distortion.

But the table showed that all Al2O3 abutments and 4 ZrO2 abutments were

fractured under static load. All Ti abutments were survived after fracture load

test.

So that, fractured specimens can be classified according to fractured

component into:

1) Favorable fracture: the fracture occurred only in the crown (all Ti

specimens and 12 ZrO2 specimens) figure (V.5).

2) Unfavorable fracture: the fracture occurred in the crown and abutment

(all Al2O3 specimens and 4 ZrO2 specimens) figure (V.6).

Favorable fracture indicates good mechanical properties of the material used

in the abutments. But unfavorable fracture indicates the material used in the

manufacturing of abutments had reasonable mechanical properties.

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Ti sp ZrO2 sp

Figure (V.5): Favorable fracture (fracture of crown only) in Ti and ZrO2

Figure (V.6): Unfavorable fracture (fracture of crown and abutment) in Al2O3

group.

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Fractured Mode in favorable fracture:

Fracture of the crowns of all specimens was approximately the same (did

not affected by abutments type). Fracture was started at midway between

central fosse and palatal cusp tip figure (V.7), and then extended mesially

and distally towards buccal side until complete fracture of crown ( fractured

started near palatal cusp tip but end near to buccal finish line) figure (V.8).

The fracture was included veneering and coping. In some specimens the

fracture was extended as two lines on mesial or distal surface, one extended

bucally and another extended palataly made triangles of fracture

Fracture mode in unfavorable fracture:

Crowns fractured in the same manner of favorable fracture. The abutments

failed in proximity to under the point in which the fractured started in crown

(in the palatal third) and the fractured continued mesialy and distally to

proximal finish line or to half of abutment figure (V.9).

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Figure (V.7): Occlusal view of fractured crown shows the fracture line

cross mesiodestally midway between central fosse and palatal cusp tip.

Figure (V.8): lateral view of fractured crown.

Figure (V.9): fracture mode of unfavorable fracture.

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SEM analysis results:

With magnification of 35 times, electron microscope showed the distortions

in the crowns in all groups were occurred in veneering and coping. Also it

showed there was no difference in fracture mode of the crown in all

specimens figure (V.10).

To detect locations and mode of failure of crown and abutment,

magnification was increased until 500 times. Electron microscope showed the

fracture was occurred intercrystally (between crystals) of press able ceramic

(crown), alumina and zirconia ceramic (abutment).

The fracture surfaces of the crown of all specimens were the same. The

surface was slightly rough due to presence of adequate number of leucite

crystals and uniform distributions of them figure (V.11A).

The fracture surface of alumina abutments was rough with distinct grains

due to presence of highly dense -alumina (Al2O3 ceramic has 3.9 g/cm3

density, 2.5 μm particle size and 5% vol porosity) figure (V.11B).

Fracture surface of zirconia abutments was moderate rough with small grain

size (5 μm in length, 0.8 μm in diameter. It has 8% vol porosity) figure

(V.11C).

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A

B

C

Figure (V.10): SEM scanning at 35 times magnification showing the fracture

of IPS crown in Al2O3, ZrO2 and Ti specimens respectively.

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Fractured surface of Y2O3- ZrO2 abutment showing its grain structure

(density 6 g/cm3, particle size 0.4 μm and 1% vol porosity).

Fractured surface of Al2O3 abutment showing its grain structure (3.9 g/cm3

density, 2.5 μm particle size and 5% vol porosity).

Fractured surface of IPS e-max crown showing highly interlocked lithium

disilicate crystals 5 am in length, and 0.8 am in diameter and 8% vol prosity

Figure (V.11): SEM scanning at 500 times magnification of ZrO2, Al2O3

abutments and IPS crown surfaces respectively.

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VI.1 Materials:

Dental implants are considered an essential treatment modality. Published

data have demonstrated high success rates for implants placed in partially

edentulous arches for the replacement of both single and multiple teeth. Also

implants decrease bone resorption and protect adjacent teeth from reduction

(Hero et al., 2004).

Dental implants and abutments are usually fabricated out of commercially

pure titanium, primarily because of it's highly corrosion resistance. Also

titanium is biocompatible and has good mechanical properties. Finally,

titanium is bioactive material (induce bone to grow around it) (Zarb et al.,

2004).

However, despite numerous modifications to the fabrication and design of

metal abutments, there is the disadvantage of metallic components showing

especially in patients with a gummy smile or a high lip line. The resultant is

dull grayish background may give the soft tissue an unnatural bluish

appearance (figure VI.1). The presence of a gray gingival discoloration may

be attributed to a thin gingival biotype that is incapable of blocking reflective

light from the metallic abutment surface. So that to achieve optimal

mucogingival esthetics, ceramic abutments were developed (Wael et al.,

2006).

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Figure (VI.1): Metal gray gingival discoloration due to use metal abutment

(Philip and James, 2003).

Zirconia (ZrO2) and Alumina (Al2O3) ceramics are used in implant

abutments recently due to many reasons:

They have nice esthetic properties to overcome gingival gray discoloration

of metallic abutment (Considering the shade of the Al2O3 ceramic, which

closely resembles that of the natural tooth, Al2O3 abutments provide certain

esthetic advantages over the more whitish ZrO2 abutments) (figure VI.2)

(Wael et al., 2006).

Figure (VI.2): A, Frontal view of cast. B, Frontal view of zirconia abutment

metal abutment (Philip and James, 2003).

ZrO2 and Al2O3 are also biocompatible to human body, non cytotoxic and

don't induce bacterial colonization on their surface (bacterial colonization on

their surface less than that on tooth surface) (Kohal et al., 2008).

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ZrO2 and Al2O3 ceramic have superior mechanical properties in

comparison with other ceramic of ceramics table (VI.1) (Kohal et al., 2008).

Type of porcelain

Point of

variation

Feldspathic ceramic

Heat pressed ceramic (IPS)

Aluminous oxide ceramic

Zirconia ceramic

Flexural strength

70MPa 400MPa 450MPa 900-1400MPa

Modulus of elasticity

69GPa 95GPa 350GPa 210GPa

Fracture toughness

1.5MPa 2.75MPa 5-6MPa 10MPa

TableVI.1: comparison of mechanical properties of alumina &zirconia with

other ceramics.

A glass-ceramic crown with low load capability (IPS Empress flexural

Strength = 400 MPa, fracture toughness 2.75 MPa) was used on the

abutments. The authors did not expect any negative effect on the fracture

toughness of the assembled specimens. The empress glass-ceramic crown

was selected for esthetic and functional reasons. Favorable clinical long-term

results were reported for IPS Empress glass-ceramic crowns on natural teeth.

Finally, it was chosen due to the feasibility of the lost-wax technique

supported by a silicone index, which reproduced the standard crowns; this

technique facilitated the fabrication of a large number of precise replicas

(Yuldrim et al., 2003).

To ensure maximum bond strength between the glass-ceramic crowns and

the Al2O3, Ti and ZrO2 abutments, the universal self adhesive resin cement

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(Rely X ARC) was used. It has high shear bond strength to glass ceramic

crown (23-25MPa) and to zirconia (25-30MPa) (Caughman et al., 2009).

The marginal integrity reaches to 99% for Rely-X which prevents the

leakage of fluids between restoration/tooth with the cement which decrease

the resorption and failure of cementation (Fraga et al., 2007).

Also it has unique chemistry raises its PH value to a neutral level of 7

quickly after setting. This contributes to the material becoming hydrophobic,

meaning it is better able to resist water uptake and more stable and durable

over time (Lu et al., 2005).

Rely-X has good mechanical properties in comparing with other cement

type. It has high flexural strength (80MPa), high compressive strength

(244MPa). Also it has surface hardness for 280MPa and high modulus of

elasticity (8.4GPa). These advanced mechanical properties lead the cement to

tolerate the forces of oral cavity (Hofmann et al., 2007).

Dimensional stability of the cement is an important consideration

especially when cementing all ceramic restorations. Rely-X cement shows

comparable and low expansion values (0.4%) that proves it to be safe for the

cementation of all-ceramic restorations (Caughman et al., 2009).

Rely-X shows less film thickness after setting (17μm), this leads the

material to give its good adhesive and mechanical properties without any

effect on the adaptation of restoration to superstructure (Jung et al., 2002).

Finally, it is available in different shades (A2 Universal, A3 Opaque and

Translucent shades) to enhance esthetic properties especially with all ceramic

restoration (Fraga et al., 2007).

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VI.2 Tests Parameter:

Before performing in vivo studies or applying materials for clinical use,

in vitro tests should be undertaken to prove materials' applicability and

performance. In vitro tests can be performed in a short period of time and

have the advantages of reproducibility and the possibility of standardizing

test parameters (Yuldrim et al., 2003).

However, each in vitro test may represent only one approach to a clinical

situation. The more closely a test simulates the clinical condition, the more

likely the results are clinically relevant (Zarb et al., 2004). It has been shown

that ceramic restorations accumulate damage during cyclic loading and

thermal cycling. The accumulated damage weakens the ceramic restoration

and can cause clinical failures (Cleries, 2009).

Intraoral occlusal forces create dynamic repetitive loading. Therefore,

instead of using monotonic static loading, it is more clinically relevant to test

the specimens under physiological fatigue load. Adding moisture and

controlled temperature to the environment was found to be important when

measuring the fracture or fatigue strength of dental ceramics (Wael et al.,

2006).

Exposure to water was been found to affect the mechanical properties of

all-ceramic restorations. Also, storage in water for extended periods has been

shown to alter the failure of all-ceramic materials and to weaken the bond

strength. Furthermore, temperature changes also lead to slow flaw

propagation. Thus, testing of all-ceramic restorations and materials should

combine these variables to more closely simulate the clinical situation (Zarb

et al., 2004).

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Some authors considered the functional forces that arise during mastication

or swallowing, which usually range between 2 to 50N. Hence, a cyclic

loading force of 10N was used to approach a clinically relevant condition. It

was shown that humans have an average of 240,000 masticators cycles per

year. Temperature of food and drinks tolerated by oral cavity between 5-

50°C. So that 600 thermal cycles between 5-50°C, and 240.000 dynamic

cyclic loading with 10N load were equaled 1 year of clinical used (Blixt et

al., 2001).

VI.3 Results:

Highest fracture load was recorded in current study for specimens restored

with titanium abutments because the titanium has superior mechanical

properties than ceramics (yield strength is 650MPa for titanium while is

450MP for zirconia), also titanium has the highest strength-to-weight ratio of

any metal (Johansson et al., 2007).

The specimens restored by ZrO2-ceramic abutments have higher fracture

loads than Al2O3 were expected due to many reasons:

Y2O3- partially-stabilized ZrO2 ceramic displays twice the flexural

strength (900 MPa to 1400 MPa) and fracture toughness (7 to 10 MPa · m1/2)

than Al2O3 ceramic that have 450MPa flexural strength and fracture

toughness value between 5 and 6 MPa ⁄ m1/2 (Kohal et al., 2008).

The enhanced strength can be also explained by micro structural

differences: Y2O3-partially-stabilized ZrO2 ceramic has a higher density (6

g/cm3) and a smaller particle size (0.4 μm) than Al2O3 ceramic (3.9 g/cm3

density and 2.5 μm particle size) (Strub and Gerds, 2003).

The metastable tetragonal crystalline structure at room temperature is

considered the main reason for the superior fracture strength of ZrO2 ceramic.

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This structure represents an efficient mechanism against flaw propagation and

has a strong impact against sub critical crack growth (Li and Duchyne,

2008).

But in current study zirconium abutment did not shown twice fracture

resistance of alumina that because many reasons: ZrO2 ceramic exhibits 1 to

10 lower thermal conductivity than Al2O3 ceramic. So temperature peaks can

alter the metastable tetragonal crystalline phase of partially stabilized ZrO2

ceramic (the transformation from tetragonal to monoclinic is completed by a

volume increase of approximately 3% to 5%. These volume changes will lead

to very high inner structure tensions and component fracture). Therefore these

abutments should be more endangered than Al2O3-ceramic abutments by heat

producing surface treatments, which produce high temperature spots because

of the very slow heat dissipation and there is controversy over whether this

would lead to a reduction in the fracture resistance of the material (Gehrke et

al., 2006).

Also zirconia is sensitive to changes in humidity and temperature. Long-

term exposure of zirconia ceramics to humidity and thermal cycling leads to a

slow, low-temperature degradation of the material that might not become

significant before several years have passed. It is well known that the

mechanical quality of machined ZrO2 ceramic is closely related to the cutting

abilities of diamond tools. This is confirmed within this study where there

was no deterioration of the Y2O3-partially-stabilized ZrO2 ceramic abutments

because of the milling process (Stuart et al., 2007).

All Al2O3 ceramic abutments were fractured with their crowns by static

load machine during this study (unfavorable fractured). This lead to increase

fracture load of Al2O3 ceramic abutments (load required to fracture crown

and abutment more than that of crown only) (Wael et al., 2006).

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Fracture of crowns was started at the interface between copings and

veneering layer, such a phenomenon can be explained by the different

coefficient of thermal expansion between coping and veneering layer.

Coefficient of thermal expansion (CTE) of veneering (18 × 10¯6/°K) is higher

than that of coping (10.5 × 10¯6/°K), after firing, during the cooling process,

the veneering is subjected to a more evident shrinkage than coping which

leads to stress concentration at this area and considered as weak component

in the ceramic restoration (Denry et al., 2010).

Unfavorable fracture (fracture of the crown and abutment in all Al2O3 and 4

ZrO2 specimens) observed in current study was occurred due to:

Flexural strength (which is a mechanical parameter for brittle material

indicate ability of these material's to resist deformation under load) of Al2O3

abutment (450MPa) is approximately near to that of IPS crown (400MPa)

(Goodacre et al., 2003).

ZrO2 abutments in comparison with Al2O3 ones have twice flexural

strength (Kohal et al., 2008), but ZrO2 abutments was more sensitive to aging

which lead to decrease their mechanical properties and may be cause fracture

of some abutments (Luthy et al., 2005).

Also ZrO2 and Al2O3 ceramic abutments are brittle material, so that any

flaws or cracks may arise naturally in a material or after aging lead to weaken

the material, and, as a result, sudden fractures can arise at stresses below the

yield stress (the stress at which the material begins to deform plastically).

Sudden, catastrophic fractures typically occur in brittle materials that don't

have the ability to plastically deform and redistribute stresses (Kenneth et al.,

2006).

Favorable fracture (in all Ti abutments and 12 ZrO2) was occurred due to

superior mechanical properties of Ti and ZrO2 than Al2O3 (fracture resistance

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of Ti is 1400MPa, 950MPa for ZrO2 while it is 450MPa for Al2O3) (Autor,

2003).

Also Ti has good thermal conductivity, so that Ti less effected by thermal

aging than ZrO2 and Al2O3 (no internal stresses occurred in Ti) and

subsequent the force need for fracture Ti abutments is higher than that for

ZrO2 and Al2O3 abutments (Craig et al., 2006).

Ti has good ductility, malleability, surface hardness and high bond

strength between its crystals, so cracks or flaws (weakening point) were not

occurred on its surface after aging and no sudden fractures can arise at

stresses below the yield stress due to Ti have the ability to plastically deform

and redistribute stresses (Stuart et al 2007).

Under SEM the fracture in the crowns and abutments was occurred inter

crystals not through the crystals because the force needs to break inter crystal

bond much less than that requires for break the crystals (Kenneth et al.,

2006).

The fracture surface of IPS crown under EM showed highly interlocked

lithium disilicate crystals, 5 μm in length, and 0.8 μm in diameter. This

interlocked microstructure and layered crystals are contributed to

strengthening. So that the crack propagation was easy along the cleavage

planes, but more difficult across the planes (Denry et al., 2010).

The fracture surface of alumina abutment under EM showed dense

alumina with distinct grains and rough fracture surface. The dense crystals

are contributed to strengthening alumina due to decrease matrix content and

shortened inter crystals bond length. So that, the fracture was occurred though

this small bond (Li and De Groot 2009).

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The fracture surface of ZrO2 abutment under EM showed small equiaxed

grains and rough fracture surface. ZrO2 fracture surface has less porosity and

micro cracks (Denry et al, 2010).

The fracture resistance of all-ceramic restorations on implants was

evaluated in other in vitro studies:

Yuldirim study in agreement with current study in the fracture load value

of ZrO2 (737.6 N) abutments was higher than that of Al2O3 (280.1N)

abutments (Yuldirim et al., 2003). But the difference in the fracture load

value between ZrO2 and Al2O3 in Yuldirim study was more than that in

current study because there was no artificial aging in Yuldirim study which

the zirconia more sensitive to it than alumina, also abutment's manufacturer

in Yuldirim study (Noble Biocare) was differed than that of current study

(Biohorizons).

Butz study was concluded that the median fracture loads were 294 N, 239

N, and 324 N for the zirconia, alumina, and titanium abutment groups which

in agreement with that in current study in arranged value (Ti had highest

fracture load value then ZrO2 and finally Al2O3). But the fracture load value

found for the abutments in Butz study are lower than those reported for

current study. This difference may be explained by methodological issues, i.e.

in Butz study the static load measurement was stopped after a deflection of 4

mm, while current study continued until a deviation from the linear slope in

the load displacement graph occurred. Also in Butz study the artificial aging

was 1.200.000 cycles but in current study 240.000 cycles (high artificial

aging cycles lead to increase damage effect on specimen then low fracture

load value require to fractured specimen). The specimens in Butz study were

restored with complete metal crowns instead of all-ceramic crowns as in

current study to obscure the cause of failure; i.e. abutment-related or crown-

related.

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Wael study found that median fracture loads were 443.6N, 422.5N, and

1454N for the zirconia, alumina, and titanium abutment groups, respectively

which is in agreement with current study in the highest fracture load value

was required for Ti abutments then ZrO2 abutments and Al2O3 abutments

were the weakest component in Wael study. But in Wael study the fracture

load value of ZrO2 abutments was near to that of Al2O3 abutments and the

both were less than that reported in current study due to the aging in Wael

study was higher than that of current study so the fracture load value required

in Wael study was lowered. Also ZrO2 abutments were more sensitive to

artificial aging than Al2O3 ones, so the fracture load value of these abutments

affected more than Al2O3 and lowered become near to that of Al2O3 ones.

Finally, the manufacturer (Nobel Biocare AB, Goteborg, Sweden) of the

abutments in Wael study was deferent than that of current study (Biohorizon,

USA).

The fracture resistance of ZrO2 abutments were evaluated and compared

with titanium ones in another vitro study (Kohal et al., 2006). In this vitro

study 16 Ti abutments restored with PFM crowns and 32 ZrO2 abutments 16

restored with Empress®-1 crown and 16 with Procera® crowns. Eight samples

of each group were exposed to a long-term load test in the artificial mouth

(1.2 million chewing cycles) then static load fracture. Another samples in

each group exposed to static loading immediately. the fracture load in the

titanium abutments–PFM crown group without artificial loading ranged

between 420 and 610 N (mean: 531.4 N), between 460 and 570 N (mean:

512.9 N) in the Empress®-1 crown group, and in the Procera® crown group

the values were between 475 and 700 N (mean: 575.7 N) when not loaded

artificially. The results when the specimens were loaded artificially with 1.2

million cycles were as follows: the titanium implant–PFM crowns fractured

between 440 and 950 N (mean: 668.6 N), the Empress®-1 crowns between

290 and 550 N (mean: 410.7 N), and the Procera® crowns between 450 and

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725 N (mean: 555.5 N). Authors said zirconia implants restored with the

Procera® crowns like titanium ones possibly fulfill the biomechanical

requirements for anterior teeth.

Kohal study disagreement with current study and all previous study in the

fracture load value of Ti abutments with PMF crowns after aging were more

than that without aging and no significant difference of aging on fracture load

value of ZrO2 abutments with Procera® crowns, this was may be due to

different abutment manufacturer or tests parameter tools.

The effect of aging on ZrO2 abutments was evaluated in study (Gehrk et

al., 2006). In this vitro study Cercon zirconium abutments without

restorations were divided into three groups. First group was exposed directly

to static loading. Second one was exposed to 10.000 cycles of artificial aging

then static loading. Third one to 5 million cycles of aging before static

loading. The fracture load values were 672N, 403N and 269N for group

respectively.

Gehrk study in agreement with current study and all previous study in

ZrO2 abutments affect by aging.

Finally, current study was on premolar which differ than all previous study

that done on central incisor (different tooth position lead to different load

direction, magnitude and different thickness of the crown).

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VII Summary

The aim of this study is to compare Zirconium &Alumina abutments versus

conventional Titanium abutments supporting all ceramic crowns in vitro by:

1) Measuring fracture resistance of crown and abutment.

2) Analyzing mode of failure using SEM.

Forty eight analogues (resembling implant) with a diameter of 4 mm and

length of 13 mm (Biohorizos, USA) represent missed root of upper first

premolar were used in this study. The implants were divided, according to the

type of abutments used into three groups (of 16 specimens each):

Group I: titanium abutments (Esthetic Titanium abutment, Biohorizons, USA) were used.

Group II: ready made zirconia abutments (ZrO2) (Esthetic Ceramic

abutment, Biohorizons, USA) were used.

Group III: ready made alumina abutments (Al2O3) (Esthetic Ceramic

abutment, Biohorizons, USA) were used.

All abutments (Ti, ZrO2 & Al2O3) had standard measurements: a deep

chamfer finish line of 1 mm depth and a total height of 6 mm &4mm

diameter.

Glass infiltrated ceramic crown system was used (IPS e-max press MO).

The specimens (analogues and abutments) were embedded in special

specimen holders using an auto polymerizing acrylic resin vertically to the

horizontal plane to simulate clinical conditions.

Relay-X ARC was used to cement crowns to abutments. Internal surface of

the crown was etched with 4.5% hydrofluoric acids for 60sec, cleaned with

water and dried with air. Abutments were sandblasted for 60 seconds. Resin

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was mixed according to the manufacturer's guidelines to lute the crown to the

abutment.

Each specimen was stored in bottle filled with distilled water for 24 hours

before application of thermal and mechanical loading.

After storage of distilled water, thermo cycling was done in thermo cycling

apparatus for 600 cycles from 5°C to 55°C with 2 min. dwell time, 10 sec.

transfer time. Then the crown of specimen was covered with load stamp and

subjected to maximum vertical load of 10 kg. With cyclic frequency of 1.7

Hz for 240.000 cycles which correspond to 12 months of clinical service.

Then the fracture resistance of the specimens was tested by computer

controlled materials testing machine (Model LRX-plus; Lloyd Instruments

Ltd., Fareham, UK).

Finally, specimens were scanned by electron (JEOL-JSM-5200LV Japan)

microscope to evaluate the failure mode at 35 times magnifications, then at

500 times magnifications.

The mean fracture loads for titanium specimens were 787.5±82.97N, where

that were 779±39.7N for ZrO2 specimens and 562.9±65.49N for Al2O3 ones.

Conclusion:

1) All of the types of implant-supported restorations tested have the

potential to withstand physiologic occlusal forces applied in the

premolar region (450N).

2) Unfavorable fracture occurred in all Al2O3 and 4 ZrO2 specimens

indicated unfavorable behavior of this material after aging.

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