Biomaterials in dental implants 12 /orthodontic courses by Indian dental academy

BIOMATERIALS INBIOMATERIALS IN
DENTAL IMPLANTSDENTAL IMPLANTS
INDIAN DENTAL ACADEMYINDIAN DENTAL ACADEMY
Leader in Continuing Dental EducationLeader in Continuing Dental Education
www.indiandentalacademy.comwww.indiandentalacademy.com

INTRODUCTION
For many years, implants of varied types have been
used in dentistry to augment or replace hard and soft
tissue components of the jaws. Currently, implant
materials include grade 2 commercially pure titanium,
titanium 6% aluminium 4% vanadium, surgical- grade
cobalt-chromium-molybdeneum, aluminium oxide in
single crystal or polycrystalline form, hydroxyapatite,
tricalcium phosphate and calcium aluminate.

The choice of material for a particular implant application
will generally be a compromise to meet many different
required properties such as mechanical strength,
machinability, elasticity, chemical properties, etc. There
is, however, one aspect that is always of prime
importance; namely, how the tissue at the implant site
responds to the biochemical disturbance that a foreign
material presents.
The most critical and debtable aspect is biocompatibility,
Dr. John Autian regards biocompatibility as that which
has no significant harm to the host.
Dr. Jonathan Black suggested that the term “biologic
performance” is more appropriate than biocompatibility
to represent the various interactions between host and
the material. www.indiandentalacademy.comwww.indiandentalacademy.com

GPT 7 defines “biocompatible” as capable of existing in
harmony with the surrounding biologic environment.
And “biomaterial” is any substance other than a drug
that can be used for any period of time as a part of a
system that treats, augments or replaces any tissue,
organ or function of the body.
Biocompatibility is dependent on the basic bulk and
surface properties of the biomaterial. All aspects of basic
manufacturing, finishing, packaging and delivering,
sterilizing, and placing (including surgical) must be
adequately controlled to ensure clean and non
traumatizing conditions.

Man has been searching for ways to replace missing
teeth for thousands of years. The first evidence of the
use of implants dates back to 600AD in the Mayan
population
•Ancient Egyptians used tooth shaped shells and ivory to
replace teeth.
•The Etruscans, living in what is now modern Italy,
replaced missing teeth with artificial teeth carved from
the bones of oxen.
•In the 1700s John Linter suggested the possibility of
transplanting teeth of one human into another
HISTORY OF MATERIALS AND DESIGNS

In 1809, Maggiolo fabricated a gold implant which was
placed into fresh extraction sockets to which he attached
a tooth after a certain healing period.
In 1886 Edmunds was the first in the US to implant a
platinum disc into the jawbone, to which a porcelain
crown was fixated.
In 1887, a physician named Harris attempted the same
procedure with a platinum post, instead of a gold post.
In the early 1990s Lambotte fabricated implants of
aluminum, silver,brass,red copper, magnesium,gold and
soft steel plated with gold and nickel.
Greenfield in 1909 made a lattice cage design of
iridoplatinum. www.indiandentalacademy.comwww.indiandentalacademy.com

•Early pioneers in this field include Dr. A.E. Strock,
who, in 1931 suggested using Vitallium r, a metal alloy,
for dental implants.
•Surgical cobalt chromium molybdenum alloy was
introduced to oral implantology in 1938 by Strock.
•In 1947, Manlio Formiggini of Italy developed an
implant made of tantalum. At the same time, Raphael
Chercheve designed a double delinked spiral implants
made of a chrome-cobalt alloy.

•By 1964, commercially pure titanium was accepted as
the material of choice for dental implants, and since that
time, almost all dental implants are made of titanium.
The body does not recognize titanium as a foreign
material, resulting in less host rejection of the implant.
In the 1960s, emphasis was placed on making the
biomaterials more inert and chemically stable within
biologic environments. The high purity ceramics of
aluminum oxide, carbon, and carbon – silicon
compounds and extra low interstitial (ELI) grade alloys
are classic examples of these trends.

In 1975 the first synthodont aluminium oxide implant
was placed in a human
Vitreous carbon implants were first placed in early
1970 by Grenoble
In early 1980s Tatum introduced Omni R implant
made of titanium alloy root form implant with horizontal
fins.
Niznick in 1980 introduced Core-vent, an endosseous
screw implant manufactured with a hydroxyapatite
coating.
Calcitek corporation began manufacturing and
marketing its synthetic polycrystalline ceramic
hydroxyapatite coated cylindrical post titanium alloy
implant.

In 1985, Straumann Company designed plasma
sprayed cylinders and screws to be inserted in a one
stage operation.
Brane mark devoted 13 years conducting animal
studies to determine the parameters under which
osseointegration would occur. Based on his study
titanium was the made the material of choice.

Selection, Evaluation and Preparation of
Biomaterials
Selection
Types : four categories
 Metals and metal alloys
 Ceramics including carbon
 Synthetic polymers
 Natural materials includes use of bone grafts for ridge
augmentation
Selection is based on:
1.The expected life time of the implant
2.Mechanical requirements

Evaluation of implant material
 Bulk characterization
 Surface characterization
Bulk material parameters important to evaluation
•Mechanical properties
•Elastic modulus
•Plastic deformation
•Tensile strength
•Fatigue
•Physical properties
•Hardness
•Thermal
•Wear
•Density
•Chemical stability
•Toxicity
•conductivity

Surface characterization
Surface properties of an implant are fundamental to the
success of the implant
Key parameters for evaluation are
•Surface energy, surface tension, chemical composition
and stability
•Morphology and texture
•Thickness of surface coating or oxide layer surface
electrical properties
•Corrosion resistance

PHYSICAL AND MECHANICAL
PROPERTIES.
Forces exerted on the implant material consist of
tensile, compressive, and shear components. As for
most materials, compressive strengths are usually
greater than their shear and tensile counterparts.
When present, parafunction (nocturnal and/or
diurnal) can be greatly detrimental to longevity because
of the mechanical properties, such as maximum yield
strength, fatigue strength, creep deformability, ductility,
and fracture. Limitations of the relevance of these
properties are mainly caused by the variable shape and
surface features of implant designs.

A different approach to match more closely the
implanted material and hard tissues properties led to the
experimentation of polymeric, carbonitic, and metallic
materials of low modulus of elasticity.
Because bone can modify its structure in response to
forces exerted on it, implant materials and designs must
be designed to account for the increased performance of
the musculature and bone in jaws restored with
implants. The upper stress limit decreases with an
increasing number of loading cycles sometimes reaching
the fatigue limit after 106 to 107 loading cycles. That is,
the higher the applied load, the higher the mechanical
stress, and the greater the possibility for exceeding the
fatigue endurance limit of the material.

In general, the fatigue limit of metallic implant materials
reaches approximately 50% of their ultimate tensile
strength. However, this relationship is only applicable to
metallic systems and polymeric systems have no lower
limit in terms of endurance fatigue strength.
Ceramic materials are weak under shear forces because
of the combination of fracture strength and no ductility,
which can lead to brittle fracture.
Metals can be heated for varying periods to influence
properties, modified by the addition of alloying elements
or altered by mechanical processing such as drawing,
swaging, or forging, followed by age or dispersion
hardening, until the strength and ductility of the
processed material are optimized for the intended
application. www.indiandentalacademy.comwww.indiandentalacademy.com

The modifying elements in metallic systems may be
metals or non metals. A general rule is that constitution
or mechanical process hardening procedures result in
an increased strength but also invariably correspond to a
loss of ductility. This is especially relevant for dental
implants.
Most all consensus standards for metals (American
Society for Testing and Material (ASTM), International
Standardization organization (ISO). American Dental
Association (ADA) require a minimum of 8% ductility to
minimize brittle fractures. Mixed microstructural phase
hardening of austenitic materials with nitrogen (e.g.
stainless steels) and the increasing purity of the alloys
seem most indicated to achieve maximum strength and
maintain this high levels of possible plastic deformation.www.indiandentalacademy.comwww.indiandentalacademy.com

7.9
170-200
240-300
600-700
200
300
35-55 www.indiandentalacademy.comwww.indiandentalacademy.com

CORROSION AND BIODEGRADATION
Corrosion is a special concern for metallic
materials in dental implantology because implants
protrude into the oral cavity where electrolyte and
oxygen compositions differ from that of tissue fluids. In
addition, the pH can vary significantly in areas below
plaque and within the oral cavity. This increases the
range of pH that implants are exposed to in the oral
cavity compared with specific sites in tissue.
Plenk and Zitter stated that galvanic corrosion can be
greater for dental implants than for orthopedic implants.

Galvanic processes depend on the passivity of oxide
layers, which are characterized by a minimal dissolution
rate and high regenerative power for metals such as
titanium. The passive layer is only a few nanometers
thick and usually made of oxides or hydroxides of the
metallic elements that have greatest affinity for oxygen.
In reactive group metals such as titanium, niobium,
circonium, tantalum, and related alloys, the base
materials determines the properties of the passive layer.
However, titanium, tantalum, and niobium oxides cover a
markedly larger zone of environmental stability
compared with chromium oxides.

There is a risk of mechanical degradation, such as
scratching or fretting of implanted materials, combined
with corrosion and release into bone and remote organs.
Lung, Willert, and Lemons, have extensively studied the
corrosion of metallic implants.
Many of the basic relationships specific to implant
corrosion have been presented by Steinemann and
Fontana and Greene.
Mears addressed concerns about galvanic corrosion and
studied the local tissue response to stainless steel and
cobalt chromium molybdenum (Co-Cr-Mo) and showed
the release of metal ions in the tissues.
Williams suggested that three types of corrosion were
most relevant to dental implants, stress corrosion
cracking, galvanic corrosion and fretting corrosion.www.indiandentalacademy.comwww.indiandentalacademy.com

Crevice corrosion
Another problem of localized corrosion of particular
importance in implant materials is crevice corrosion. This
occurs when a crevice is formed by covering or shielding
a portion of the metal from the corrosive medium. The
area between a metal post and a prosthetic tooth is one
eg. The figure shows an idealized crevice and the
surrounding environment. The shielded area has limited
access to the surrounding solution which contains
corrosive species such as Cl ions. Since the access is
limited ,metal ions and hydrogen ions build up with a
corresponding lowering of the oxygen concentration.
The Cl ions move into the crevice due to charge effects
and cause more damage.

The shielded area becomes the anode and the non
shielded becomes the cathode. The lack of oxygen in
the crevice environment as well as the pH and Cl ion
content act as crucial factors in creating corroding
crevice.

STRESS CORROSION CRACKING
The combination of high magnitudes of applied
mechanical stress plus simultaneous exposure to a
corrosive environment can result in the failure of metallic
materials by cracking, where neither condition alone
would cause the failure. Williams presented this
phenomenon of stress corrosion cracking (SCC) in
multicomponent orthopedic implants.
Lemons and others hypothesized that it may be response
for some implant failures in view of high concentrations of
forces in the areas of the abutment to implant body
interface.

Stress corrosion

Most traditional implant body designs under three
dimensional finite element stress analysis show a
concentration of stresses at the crest of the bone
support and cervical one third of the implant. This tends
to supports potential SCC at the implant interface area
(i.e., a transition zone for altered chemical and
mechanical environmental conditions). This has also
been described in terms of corrosion fatigue (i.e, cyclic
load cycle failures accelerated by locally aggressive
medium). In addition, non passive prosthetic super
structures may in corporate permanent stress, which
strongly influences this phenomenon under loaded
prostheses.

Galvanic corrosion (GG) occurs when two dissimilar
metallic materials are in contact and are within an
electrolyte resulting in current to flow between the two.
The metallic materials with the dissimilar potentials can
have their corrosion currents altered, thereby resulting in
a greater corrosion rate. Fretting corrosion (FC) occurs
when there is a micromotion and rubbing contact within
a corrosive environment (such as the perforation of the
passive layers and shear directed loading along
adjacent contacting surfaces). The loss of any protective
film can results in the acceleration of metallic ion loss.
FC has been shown to occur along implant
body/abutment/superstructure interfaces.

Normally, the passive oxide layers on metallic substrates
dissolve at such slower rates that the resultant loss of
mass is of no mechanical consequences to the implant.
A more critical problem is irreversible local perforation of
the passive layer that is often caused by chloride ions,
which may result in localized pitting corrosion. Such
perforations can often be observed for iron chromium
nickel – molybdenum (Fe-Cr-Ni-Mo) steels that contain
an insufficient amount of the alloying elements
stabilizing the passive layer (Cr and Mo) or local regions
of implants that are subjected abnormal environments.
Even ceramic oxide materials are not fully degradation
resistant.

Pitting corrosion
Galvanic corrosion

Corrosion like behavior of ceramic materials can then be
compared with the chemical dissolution of the oxides
substrates. An example of this is the solubility of
aluminum oxide as alumina or titanium oxide as titania.
Most metallic oxides and non metallic substrates have
amorphous – hydroxide inclusive structures, whereas
bulk ceramics are mostly crystalline. The corrosion
resistance of synthetic polymers depends not only on
their composition and structural form but also on the
degree of polymerization. Unlike metallic and ceramic
materials, synthetic polymers are not only dissolved but
also penetrated by water and substances from biologic
environments.

Galvanic attack occurs when two dissimilar metals
touch in an electrolyte solution
Pitting corrosion occurs at a specific location due to
chemical breakdown, perforation or penetration of
passive film
Crevice corrosion due to lack of oxygen at the site of
corrosion
Stresses and stress corrosion with emphasis on
elimination of possible prestressing implants.
Corrosion fatigue was described in connection with
cyclic stresses applied to implants.
Fretting corrosion, which is a result of abrasion and
produces debris.

TOXICITY AND CONSIDERATION
Toxicity is related to primary biodegradation
products (simple and complex cations and anions),
particularly those of higher atomic weight metals.
Factors to be considered include (1) the amount
dissolved by biodegradation per time unit, (2) the amount
of material removed by metabolic activity in the same
time unit, and (3) quantities of solid particles and ions
deposited in the tissue and any associated transfers to
the systemic system.
The toxicity is related to the content of materials toxic
elements and that they may have a modifying effect on
corrosion rate.

The transformation of harmful primary products is
dependent on their level of solubility and transfer. It is
known that chromium and titanium ions react locally at
low concentrations, whereas Co, Mo or Ni can remain
dissolved at higher relative concentrations and thus
may be transported and circulated in body fluids.
Lemons et al. reported on the formation of
electrochemical couples as a result of oral implant and
restorative procedures and stressed importance of
selecting compatible metals to be placed in direct
contact with one another in the oral cavity to avoid the
formation of adverse electrochemical couples.

The electrochemical behavior of implanted materials has
been instrumental in assessing their biocompatibility.
Zitter et al. have shown that anodic oxidation and
cathodic reduction take place in different spaces but
must always balance out through charge transfer. This
has been shown to impair both cell growth and the
transmission of stimuli from one cell to another.
Therefore an anodic corrosion site can be influenced by
ion transfer but also by other possibly detrimental
oxidation phenomena.

Charge transfer appears to be a significant factor
specific to the biocompatibility of metallic biomaterials.
Passive layers along the surfaces of titanium, niobium,
zirconium, and tantalum increase resistance to charge
transfer processes by isolating the substrate from the
electrolyte, in addition to providing a higher resistance to
ion transfers. On the other hand, metals based on iron,
nickel, or cobalt is not as resistant to transfers through
the oxide like passive surface zones.

CLASSIFICATION OF
BIOMATERIALS
METALS AND ALLOYS
 Titanium and Titanium –6 Aluminum-4 Vanadium
(Ti-6AI- 4V) and cp Ti
 Cobalt-Chromium-Molybdenum-Based Alloy
 Iron-Chromium-Nickel-Based Alloys
 Other metals and Alloys
CERAMICS
• Aluminum, Titanium and Zirconium oxide
• Bioactive and biodegradable ceramics
CARBON
• Carbon and carbon siliconwww.indiandentalacademy.comwww.indiandentalacademy.com

POLYMERS AND COMPOSITES
 Polymethylmethacrylate (PMMA)
 Polyethylene (UHMW-PE)
 Polytetrafluoroethylene (PTFE)
 Silicone rubber
 Polysulfone

DIFFERENT CLASSES OF SOLID
MATERIALS
Almost all inorganic materials that are of any
interest as construction materials consist of very dense
arrangement of their constituent atoms. They are
penetrable (often very slowly) only by diffusion of single
atoms, but do not allow passage of even the smallest
molecules.
Most of these materials are crystalline and are composed
of a large number of small crystallites. Each crystallite is
an ordered arrangement of atoms. Such materials are
called polycrystalline. Most metals and many ceramics
are polycrystalline.

In some materials the atoms are arranged in a less
ordered way, almost as in a liquid but with much less
mobility. Such materials are called amorphous. Most
important are glasses.
Many materials can take different crystalline forms in
different situations. One well known example is carbon,
which can be completely crystalline as in a diamond.
Graphite, another form of carbon, is also crystalline.
Carbon can also be amorphous. These different forms
have very different properties, which originate from their
differences in atomic arrangements.

Metals are special among the construction materials.
They are single element materials (composed of one
kind of atom), many are easily machined, they are
ductile, and they have advantageous mechanical
properties. Metals, however, are also reactive (except
the noble metals Au, Pt, Pd, etc.) and therefore usually
exist in nature as chemical compounds. One important
consequence of this reactivity is that most pure metals
are covered by an oxide layer.
Sometimes two or more different metals are mixed
in order to make better certain properties. Such metallic
mixtures are called alloys. Well known examples are
brass (63% Cu, 27% Zn) and stainless steel (Fe plus
small amounts of other metals such as Cr, Ni, V, Mo).

Many nonmetallic materials are formed as chemical
compounds between metals and other elements such as
oxides, nitrides, and carbides. Many of these materials
are classified as ceramics. Example include aluminum
oxide (AI2O3), titanium oxide and titanium nitride, and
tungsten carbide. Characteristics properties of ceramics
are their great hardness (but usually high brittleness),
good high temperature properties and chemical
inertness. Usually, they are mechanically not as strong
and advantageous as metals and they are much more
difficult to machine.

Glasses are materials related to the ceramic materials
(or they may be regarded as a particular class of
ceramics) but have an amorphous structure. Glasses
are often compounds of several elements and can
usually be formed to particular geometric structures via
their molten state or by machining.
Metals are the most versatile in organic materials in view
of their high strength and ductility, elasticity and
machinability, but sometimes it is advantageous to
combine these properties with some of the superior
properties of ceramics, for example. This combination
has led to the surface coating techniques, which
combine the best characteristics of two or more different
materials. www.indiandentalacademy.comwww.indiandentalacademy.com

For example, the mechanical strength maybe obtained
from a bulk metal whereas the corrosion or wear
resistance is obtained from a layer of ceramic material.
Metals are, in this respect, very special because they
offer this kind of combination of properties. Stainless
steel, for example, has enormous versatility due to its
bulk metallic properties, but its corrosion resistance is
the result of the very dense and chemically inert oxide
(i.e. ceramic) of 5 nm thickness that automatically forms
on the surface of this alloy upon exposure to air.
Independent of which material is chosen as an implant
material, it will be its surface that comes into contact with
the host tissue.

Titanium and Titanium –6 Aluminum-4
Vanadium (Ti-6AI- 4V)
Titanium was selected as the material of choice
because of its inert and biocompatible nature paired
with excellent resistance to corrosion.
This reactive group of metals and alloys form
tenacious oxides in air or oxygenated solutions. Titanium
(Ti) oxidizes (passivates) upon contact with room
temperature air and normal tissue fluids. This reactivity
is favorable for dental implant devices.
In the absence of interfacial motion or adverse
environmental conditions, this passivated (oxidized)
surface condition minimizes biocorrosion phenomena.

An oxide layer 10A thick forms on the cut surfaces of
pure titanium within a millisecond. Thus any scratch or
nick in the oxide coating is essentially self healing.
Titanium is further passivated by placement in a bath of
nitric acid to form a thick, durable oxide coating.
The high biocompatibility of titanium as an implant
material is connected with the properties of its surface
oxide. In air or water titanium quickly forms an oxide
thickness of 3 to 5 nm at room temperature.
Pure titanium contains 0.5% oxygen and minor amounts
of impurities such as nitrogen, carbon and hydrogen. In
its most common alloyed form, it contains 90%wt
titanium, 6%wt aluminum, 4%wt vanadium.www.indiandentalacademy.comwww.indiandentalacademy.com

Titanium can form several oxides of different
stoichiometry – TiO, Ti2O3, TiO2 – of which TiO2 is the
most common. TiO2 can have three different crystal
structures – rutile, anatase, and brookite – but also can
be amorphous.
TiO2 is very resistant against chemical attack, which
makes titanium one of the most corrosion resistant
metals, particularly in the chemical environment . This is
one contributing factor to its high biocompatibility. This
property is also shared with several other metals such
as Al which forms AI2O3 and Zr which forms ZrO2 on
their surfaces.

 Another physical property that is unique for TiO2 is its
high dielectric constant, which ranges from 50 to 170
depending on crystal structure. This high dielectric
constant would result in considerably stronger van der
Waal’s bonds on TiO2 than on other oxides, a fact that
may be important for the interface biochemistry.
 TiO2, like many other transition metal oxides, is
catalytically active for a number of inorganic and organic
chemical reactions, which also may influence the
interface chemistry.

Original Branemark
fixture
Titanium screw
Cp Ti screw
implant

 Titanium shows a relatively low modulus of
elasticity and tensile strength when compared with
most other alloys.
The strength values for the wrought soft and ductile
metallurgic condition (normal root forms and plate form
implants) are approximately 1.5 times greater than the
strength of compact bone.
In most designs where the bulk dimensions and shapes
are simple, strength of this magnitude is adequate.
Because fatigue strengths are normally 50% weaker or
less than the corresponding tensile strengths, implant
design criteria are decidedly important.

Sharp corners or thin sections must be avoided for
regions loaded under tension or shear conditions. The
modulus of elasticity of titanium is 5 times greater than
that of compact bone, and this properly places emphasis
on the importance of design in the proper distribution of
mechanical stress transfer. In this regard, surface areas
that are loaded in compression have been maximized for
some of the newer implant designs.
 Four grades of unalloyed Ti and Ti alloy are the most
popular. Their ultimate strength and endurance limit vary
as a function of their composition.

930
860
113
Ti-6Al-4V

 The alloy of titanium most often used is titanium
aluminum-vanadium. The wrought alloy condition is
approximately 6 times stronger than compact bone and
thereby affords more opportunities for designs with
thinner sections (e.g., plateaus, thin interconnecting
regions, implant-to-abutment connection screw housing,
irregular scaffolds, and porosities). The modulus of
elasticity of the alloy is slightly greater than that of
titanium, being about 5.6 times that of compact bone.
The alloy and the primary element (Ti) both have
titanium oxide (passivated) surfaces.

 Electrochemically, Ti and Ti alloy are slightly
different with regard to electromotive and galvanic
potentials when compared with other electrically
conductive dental materials. In general, titanium and
cobalt-based systems are electrochemically similar;
however, comparative elements imitating the conditions
in an aeration cell revealed that the current flow in Ti and
Ti alloys is several orders of magnitude lower than that
in Fe-Cr-Ni-Mo steels or Co-Cr alloys.
Gold, platinum, and palladium-based systems have
been shown to be noble, and nickel, iron, copper, and
silver-based systems are significantly different (subject
to galvanic coupling and preferential in vivo corrosion).

 Mechanically, Ti is much more ductile (bendable)
than Ti-alloy. This feature has been a very favourable
aspect related to the use of titanium for endosteal plate
form devices. The need for adjustment or bending to
provide parallel abutments for prosthetic treatments has
caused manufacturers to optimize microstructures and
residual strain conditions. Coining, stamping, or forging
followed by controlled annealing heat treatments are
routinely used during metallurgic processing.

 However, if an implant abutment is bent at the time of
implantation, the metal is strained locally at the neck
region (bent) and the local strain is both cumulative and
dependent on the total amount of deformation
introduced during the procedure.
This is one reason, other than prior loading fatigue
cycling, why reuse of implants is not recommended.
Also, sometimes mechanical processes can significantly
alter or contaminate implant surfaces.
Any residues of surface changes must be removed
before implantation to ensure mechanically and
chemically clean conditions.

Preparation of titanium dental implants
The nature of the surface oxide on titanium (or
any other metal) implants depends crucially on the
conditions during the oxidation and the subsequent
treatment of the implant. Preparation methods for the
dental implants used by Branemark as reported by
Adell et al., discusses how the various preparation
steps may influence the implant surface. The implants
are made from pure titanium that is shaped by carefully
controlled machining (lathing, threading, milling, etc.)
During the machining procedure, the fresh metals is
exposed to air (and lubricants or coolants) and oxidizes
rapidly. The nature of the surface oxide will depend on
the machining conditions (e.g. pressure and speed).

During the subsequent preparation steps (ultra sonic
cleaning and sterilizing) the initial surface oxide will be
modified.
Especially during the sterilizing procedure (autoclaving)
the oxide will undergo a slight growth in the elevated
temperature and humid atmosphere. Autoclaving also
might cause incorporation of OH radicals in the surface
oxide.

Spectroscopic characterization and elemental
composition of titanium implant surfaces.
There are several chemical elements present on the
oxidized titanium surface that are absent on the
reference TiO2 sample. A large carbon signal (- 40
atomic %) is always observed, as well as a smaller
nitrogen. Lower concentrations of chlorine, sulphur, and
calcium are often detected. These impurities except Ca
are confined to the outermost atomic layer, which
means that their total concentrations are in the range of
0.001 – 0.01 ug per square centimeter implant surface.
Ca, however, is found throughout the oxide layer.

The origin of these very small concentrations of
contaminants is probably adsorption of C, N, S, and Cl
containing molecules on the oxide surface during the
preparation procedures. They can easily be removed by
a slight ion etching in vacuum,
Another type of analysis indicated that the oxide also
contains relatively large amounts of hydrogen, probably
bound as OH.
Because the role of even small amounts of contaminants
on the biocompatibility of implant materials is not well
known, it is advisable to keep a high standard on the
cleaning procedures.

One recent example illustrates how an impurity of very
low concentration can dramatically change the
properties of the surface oxide. Via the textile cloths
wrapped around the container box for the titanium
fixtures during autoclaving, a very minute amount of
fluorine was deposited on the titanium surfaces. On the
most exposed parts this resulted in the growth of more
than 700-A-thick oxide films, which is more than ten
times the thickness usually found after autoclaving. The
fluorine ions obviously accelerated the oxide growth
considerably. Since the acceptance or non acceptance
of such changes by the body tissue are unknown, great
care must be taken to avoid impurities. Particular
attention should be paid to catalytically active elements,
which can profoundly influence the chemical interface
processes even at extremely low concentrations.www.indiandentalacademy.comwww.indiandentalacademy.com

Alternative surface preparation methods.
Although the present preparation procedures for
dental titanium implants have been highly successful, it
is unlikely that they are optimal from a biocompatibility
point of view. It may therefore be desirable that new
techniques are applied, by which the surface properties
of titanium (or other metals) implants can be varied in a
more controlled manner.
There exists today a large number of different
methods for more or less sophisticated surface
treatment, including anodic oxidation, plasma oxidation,
plasma cleaning, and vapor deposition.

Anodic oxidation
Anodic oxidation is an electrochemical method of
treatment. The sample to be treated is made an anode
in an electrolytic bath, and when a potential is applied
on the sample, a current will flow through the electrolyte
due to ion transport. The transport of oxygen ions
through the electrolyte builds up a passivating oxide
layer on the surface of the sample. The thickness of the
surface oxide formed depends, often linearly, on the
applied potential. Anodic oxidation thus offers a
possibility to control the thickness of the surface oxide
in a much wider range than thermal oxidation allows.

By a proper choice of electrolytes, the chemical
composition of the oxide can, to some extent, be
controlled, for example, by incorporation of mineral
ions. The crystal structure of the oxide can also be
varied by using electrolyte, current density, and oxide
thickness as parameters.
Plasma oxidation
In plasma oxidation, an oxygen plasma is used
instead of a liquid electrolyte. Plasma oxidation offers
essentially the same possibilities to control the surface
oxide but is basically a cleaner method than anodic
oxidation. Plasma cleaning is technically identical to
plasma oxidation, but used in order to increase the
surface cleanliness, which usually results in an
increase in the surface energy.www.indiandentalacademy.comwww.indiandentalacademy.com

Vapor deposition
Vapor deposition can be used to deposit desirable
atoms or continuous films on surfaces. As the name
implies, the method is based on the principle that the
material to be deposited is heated until it evaporates.
Alternatively, energetic ions can be used to vaporize the
material. The vapor is then allowed to condense on the
material to be covered. These techniques are often
referred to as physical vapor deposition (PVD).
Deposition can also be made by chemical reactions and
is then called chemical vapor deposition (CVD). With
PVD and CVD a wide range of composite materials and
surface coatings can be produced.

Future development
It is likely that these techniques will play an
increasingly important role in the future development of
implant materials. One can safely say that the limitation
lies in the methods by which biocompatibility can be
“measured”. The available tests that can decide
whether one implant material is better than the other
are inexact and time consuming. There is thus a great
need for a combination of biochemical and medical
tests that can specify relevant biocompatibility
parameters. Once such tests are available, the state of
the art of surface preparation and characterization
techniques can be combined to tailor make implant
surfaces for optimal biocompatibility.

Cobalt-Chromium-Molybdenum-Based Alloy
 The cobalt-based alloys are most often used as cast
or cast-and-annealed metallurgic condition. This permits
the fabrication of implants as custom designs such as
subperiosteal frames.
The elemental composition of this alloy includes 63%
cobalt, 30% chromium, and 5% molybdenum as the
major elements.
Cobalt provides the continuous phase for basic
properties; secondary phases based on cobalt,
chromium, molybdenum, nickel, and carbon provide
strength (4 times that of compact bone) and surface
abrasion resistance,
chromium provides corrosion resistance through the
oxide surface; www.indiandentalacademy.comwww.indiandentalacademy.com

while molybdenum provides strength and bulk
corrosion resistance.
All of these elements are critical, as is their
concentration, which emphasizes the importance of
controlled casting and fabrication technologies. Also
included in this alloy are minor concentrations of nickel,
manganese, and carbon.
Nickel has been identified in biocorrosion products,
and carbon must be precisely controlled to maintain
mechanical properties such as ductility.
Surgical alloys of cobalt are not the same as those used
for partial dentures, and substitutions should be avoided.
These alloys posses outstanding resistance to corrosion
and they have a high modulus.

In general, the as-cast cobalt alloys are the least
ductile of the alloy systems used for dental surgical
implants, and bending of finished implants should be
avoided.
 Because of the requirements of low cost and long
termclinical success these alloys have been used
extensively in many areas of surgery and dentistry.

Iron-Chromium-Nickel-Based Alloys
 The surgical stainless steel alloys (e.g., 316 Low
carbon) have a long history of use for orthopedic and
dental implant devices.
 This alloy, as with titanium systems, is used most
often in a wrought and heat-treated metallurgic
condition, which results in a high-strength and high-
ductility alloy.
 The ramus blade, ramus frame, stabilizer pins (old)
and some mucosal inert systems have been made from
the iron-based alloy.
 The ASTM F4 specification for surface passivation
was first written and applied to the stainless steel alloys.
This was done to maximize corrosion-biocorrosion
resistance. www.indiandentalacademy.comwww.indiandentalacademy.com

Of the implant alloys, this alloy is most subject to
crevice and pitting biocorrosion, and care must be taken
to use and retain the passivated (oxide) surface
condition.
Because this alloy contains nickel as a major element,
use in patients allergic or hypersensitive to nickel should
be avoided.
Also, if a stainless steel implant is modified before
surgery, recommended procedures call for repassivation
to obtain an oxidized (passivated) surface condition to
minimize in vivo biodegradation.
The iron-based alloys have galvanic potentials and
corrosion characteristics that could result in concerns
about galvanic coupling and biocorrosion if
interconnected with titanium, cobalt, zirconium, or
carbon implant biomaterials.www.indiandentalacademy.comwww.indiandentalacademy.com

In some clinical conditions, more than one alloy may be
present within the same dental arch of a patient.
For example, if a bridge of a noble or a base metal alloy
touches the abutment heads of a stainless steel and
titanium implant simultaneously, an electrical circuit
would be formed through the tissues. If used
independently, where the alloys are not in contact or not
electrically interconnected, the galvanic couple would
not exist, and each device could function independently.
 Long-term device retrievals have demonstrated that,
when used properly, the alloy can function without
significant in vivo breakdown.
 Clearly, the mechanical properties and cost
characteristics of this alloy offer advantages with respect
to clinical applications.www.indiandentalacademy.comwww.indiandentalacademy.com

Other Metals and Alloys
Many other metals and alloys have been used for
dental implant device fabrication. Early spirals and
cages included tantalum, platinum, iridium, gold,
palladium, and alloys of these metals.
More recently, devices made from zirconium, hafnium,
and tungsten have been evaluated.
Gold, platinum, and palladium are metals of relatively
low strength, which places limits on implant design.

These metals, especially gold because of nobility and
availability, continue to be used as surgical implant
materials.

CERAMICS
 Ceramics are inorganic, nonmetallic, nonpolymetric
materials manufactured by compacting and sintering at
elevated temperatures.
 They can be divided into metallic oxides or other
compounds.
 Oxide ceramics were introduced for surgical implant
devices because of their inertness to biodegradation,
high strength, physical characteristics such as color and
minimal thermal and electrical conductivity, and a wide
range of material specific elastic properties.
 In many cases, however, the low ductility or inherent
brittleness has resulted in limitations.
 Ceramics have been used in bulk forms and more
recently as coatings on metals and alloys.www.indiandentalacademy.comwww.indiandentalacademy.com

Aluminum, Titanium And Zirconium Oxides
Ceramics from aluminum, titanium, and zirconium
oxides have been used for root form, endosteal plate
form, and pin-type dental implants.
 The compressive, tensile, and bending strengths
exceed the strength of compact bone by 3 to 5 times.
 The aluminum, titanium and zirconium oxide ceramics
have a clear, white, cream or light grey color, which is
beneficial for applications such as anterior root form
devices.
 Minimal thermal and electrical conductivity, minimal
biodegradation, and minimal reactions with bone, soft
tissue, and the oral environment are also recognized as
beneficial when compared with other types of synthetic
biomaterials.

 In early studies of dental and orthopedic devices in
laboratory animals and humans, ceramics have
exhibited direct interfaces with bone, similar to an
osseointegrated condition with titanium.
 Although the ceramics are chemically inert, care must
be taken in the handling and placement of these
biomaterials. Exposure to steam sterilization results in a
measurable decrease in strength for some ceramics;
scratches or notches may introduce fracture-initiation
sites; chemical solutions may leave residues; and the
hard and sometimes rough surfaces may readily abrade
other materials thereby leaving a residue on contact.
Dry heat sterilization within a clean and dry atmosphere
is recommended for most ceramics.

Bioceram single crystal
sapphire implant
Synthodont aluminum oxide
implant

Although initial testing showed adequate mechanical
strengths for these polycrystalline alumina materials, the
long-term clinical results clearly demonstrated a
functional design-related and material-related limitation.
 The established chemical biocompatibilities, improved
strength and roughness capabilities of sapphire and
zirconia, and the basic property characteristics of high
ceramics continue to make them excellent candidates
for dental implants.

Carbon and Carbon Silicon Compounds
Carbon compounds are often classified as ceramics
because of their chemical inertness and absence of
ductility; however, they are conductors of heat and
electricity.
Extensive applications for cardiovascular devices,
excellent biocompatibility profiles, and moduli of
elasticity close to that of bone have resulted in clinical
trials of these compounds in dental and orthopedic
prostheses.
One two-stage root replacement system (Vitredent)
was quite popular in the early 1970s. However, a
combination of design, material, and application
limitations resulted in a significant number of clinical
failures and the subsequent withdrawal of this device
from clinical use. www.indiandentalacademy.comwww.indiandentalacademy.com

Advantages are
•tissue attachment; regions that serve as barriers to
elemental transfer, heat, or electrical current flow;
•control of color;
•and opportunities for the attachment of active
bimolecular or synthetic compounds.
Limitations relate
•to mechanical strength properties along the substrate-
to-coating interface;
•biodegradation that could adversely influence tissue
stabilities;
•time-dependent changes in physical characteristics;
minimal resistance to scratching or scraping procedures
associated with oral hygiene;
•and susceptibility to standard handling, sterilizing, or
placing methodologies.www.indiandentalacademy.comwww.indiandentalacademy.com

Vitreous carbon implants
In the early 1970s, with the aid of advanced materials,
Grenoble and coworkers introduced vitreous carbon
implants.
Vitreous carbon is a 99.99 % pure form of carbon with
a compressive strength of 50,000 to 100,000 pounds per
square inch,
a transverse strength of 10,000 to 30,000 psi
and a modulus of elasticity between 3 and 4 x 106 psi.
This modulus is similar to that of dentin, this is a
significant factor in reducing shearing forces at the
implant bone interface.
This implant is formed by molding resin into the
implant shape, heat treating it under nitrogen and then
vacuumizing it to evaporate the nitrogen, oxygen,
hydrogen and any impurities included in the resin.www.indiandentalacademy.comwww.indiandentalacademy.com

Vitreous carbon
Pyrolytic carbon

Pyrolytic carbon implants
Since vitreous carbon is a brittle material with limited
strength, it was not feasible to fabricate a satisfactory
vitreous carbon in the blade shape configuration.
Hence the pyrolytic carbon or LTI (low temperature
isotropic carbon) are formed in a fluidized bed by the
pyrolysis of a gaseous hydrocarbon depositing carbon
onto a preformed substrate such as polycrystalline
graphite. The silicon variety of pyrolytic carbon is
prepared by codepositing silicon with carbon to produce
stronger implant material.
The strength and its ability to absorb energy on impact
is nearly 4 times greater than that of glassy or vitreous
carbon. The modulus of elasticity of all isotropic carbon
materials is 3 to 4 x 106 psi almost similar to that of
done. www.indiandentalacademy.comwww.indiandentalacademy.com

Therefore carbon implant can bend and displace as if it
were cortical bone, thus minimizing stress
concentrations that could otherwise cause bone
resorption and implant loosening.
Of all materials carbon is the most biocompatible. The
biocompatibilty of silicon – alloyed pyrolytic carbon with
blood, soft and hard tissues is superior to that of all
other known materials. LTI carbon can interface with
blood without producing the clotting effect seen with
most other foreign materials.www.indiandentalacademy.comwww.indiandentalacademy.com

POLYMERS
The use of synthetic polymers and composites
continues to expand for biomaterial applications.
Fiber-reinforced polymers offer advantages in that they
can be designed to match tissue properties, can be
anisotropic with respect to mechanical characteristics,
can be coated for attachment to tissues, and can be
fabricated at relatively low cost.
Structural biomedical polymers
•Polymers have lower strengths and elastic moduli and
higher elongations to fracture compared with other
classes of biomaterials.

•They are thermal and electrical insulators and when
constituted as a high molecular weight system without
plasticizers, they are relatively resistant to
biodegradation.
•When compared with bone they have lower elastic
moduli with magnitudes closer to soft tissues.
• Polymers have been fabricated in porus and solid
forms for tissue attachment, replacement and
augmentation and as coatings for force transfer to soft
tissue and hard tissue regions.
• Cold flow characteristics and creep and fatigue
strengths are relatively low for SR and Pmma polymers.
• In contrast, extremely tough and fatigue cycle resistant
for PP, UHMW-PE, PTFE and afford for mechanical
force transfer. www.indiandentalacademy.comwww.indiandentalacademy.com

•Most uses have been for internal force distribution
connectors for osseointegrated implants where the
connector is intended to better simulate biomechanical
conditions for normal tooth functions.
• Indications for PTFE have grown in the last decade
because of the development of membranes for GTR
techniques. However they have a low resistance to
contact abrasion and wear phenomenon.

COMPOSITES
•Combinations of polymers and other categories of
synthetic biomaterials
•Inert polymers are combined with particulate or fibers
of carbon, aluminum oxide, hydroxyapatite and glass
ceramics.
•Some are porous while others are constituted as solid
composite structural forms.
•Biodegradable polymers like polyvinyl alcohol (PVA),
polylactides or glycolides, cyanoacrylates or other
hydratable forms have been combined with
biodegradable CaPO4 particulate or fibers.
•Used as structural scaffolds, plates, screws. Also used
as bone augmentation and periimplant defect repairs.

• Polymers and composites of polymers are especially
sensitive to sterilization and handling techniques. When
used as implants they cannot be sterilized by steam or
ethylene oxide.
• They have electrostatic surface properties and tend to
gather dust or other particulate if exposed to semiclean
air environments.
• Can be shaped by cutting or auto polymerizing in vivo
therefore extreme care must be taken to maintain quality
surface conditions of the implant.
• Porous polymers can be deformed by elastic
deformation and cleaning of contaminated porous
polymers is not possible without laboratory environment.
• Long term experience, excellent biocompatibility, ability
to control properties through composite structures.

Inserts and Intramobile elements
• Relatively low modulus of elasticity when compared
with metals and ceramics.
• High elongation to fracture, inherent toughness have
resulted in use of these polymers as connectors or
interpositional spacers for dental implants.
• Limitation being cyclic load creep and fatigue
phenomena.
• Plastic deformation and fracture.
• Inadequate long term performance and high time and
cost associated to maintenance have resulted in failures.

SURFACE CHARACTERIZATION AND TISSUE
INTEGRATION
Metal and Alloy Surfaces
Standard grades of alpha (unalloyed) titanium and alpha
beta and beta-base alloys of titanium (Ti) exist with an
oxide surface at normal temperatures, with ambient air
or normal physiologic environments that act as oxidizing
media. There is a formation of a thin oxide via
dissociation of and reactions with oxygen or other
mechanisms such as oxygen or metal ion diffusion from
and to the metallic surface, especially for titanium. This
thin layer of amorphous oxide will rapidly reform if
removed mechanically. Surface properties are due to
this oxide layer and differ fundamentally from the
metallic substrate. www.indiandentalacademy.comwww.indiandentalacademy.com

Therefore the oxidation parameters such as
temperature, type and concentration of the oxidizing
elements, and eventual contaminants all influence the
physical and chemical properties of the final implant
product. The type of oxide on surgical implants is
primarily amorphous in atomic structure (Brookite) if
formed in normal temperature air or tissue fluid
environments and is usually very adherent and thin in
thickness dimensions (less than 20 nanometers). In
contrast, if unalloyed titanium (alpha) substrates
(titanium grades 1 to 4) are processed at elevated
temperatures (above approximately 3500 C) or anodized
in organic acids at higher voltages (above 200 mV), the
oxide forms a crystalline atomic structure (Rutile or
Aanatase) and can be 10-100 times thicker.www.indiandentalacademy.comwww.indiandentalacademy.com

 Porosity, density, and general homogeneity of the
substrate are all related to this process.
 Low temperature thermal oxides are relatively
homogeneous and dense; with increasing temperatures
they become more heterogeneous and more likely to
exhibit porosity as scale formations and some have
glasslike surface oxide conditions (semicrystalline).
 Depending on the mechanical aspects of cleaning and
passivating, these amorphous or crystalline oxides can
exhibit microscopically smooth or rough topographies at
the micrometer level. However, surface macroscopic
roughness is normally introduced into the substrate
beneath the oxide zone by mechanical (grinding),
particulate blasting (resorbable blast media or other), or
chemical (acid etching) procedures.www.indiandentalacademy.comwww.indiandentalacademy.com

Tissue Interactions
 Oxide modification during in vivo exposure has been
shown to result in increased titanium oxide layer
thickness of up to 200 nm.
 The highest oxide growth area corresponded to a
bone marrow site while the lowest growth was
associated with titanium in contact with cortical regions
of bone. Increased levels of calcium and phosphorus
were found in the oxide surface layers and seemed to
indicate an active exchange of ions at the interface.
 Hydrogen peroxide environmental condition has been
shown to interact with Ti and form a complex gel. “Ti gel
conditions” are credited with attractive in vitro properties
such as low apparent toxicity, inflammation, bone
modeling, and bactericidal characteristics.www.indiandentalacademy.comwww.indiandentalacademy.com

The surface bio-interaction processes may be slow or
activated by local reactions and may cause ion release
and oxide alteration of the substrate. Local and
systemic increases of the ion concentration have been
reported. In vitro studies showed that both Ti and Ti
alloy were released in measurable quantities of the
substrate elements at the surface. Especially high rates
of ion release were observed in
ethylenediamineteraacetic acid (EDTA) and sodium
citrate solutions and varied as a function of the corroding
medium.

Integration with Titanium and Alloys
Although titanium is known to exhibit better
corrosion resistance, independent of the surface
preparation, in vivo and in vitro studies have shown that
titanium may interact with the recipient living tissues
over several years. This interaction results in the release
of small quantities of corrosion products even though
there is a thermodynamically stable oxide film.
Several studies have concentrated on the
behavior of Ti and Ti alloys in simulated biologic
environments. Williams cautioned that although titanium
can demonstrate excellent properties of its tenacious
oxide film, it is usually not sufficiently stable to prevent
“wear and galling” in bearing systems under load.

Titanium implants may be etched with a solution of nitric
and hydrofluoric acids to chemically alter the surface
and eliminate some types of contaminant products. The
acids very rapidly attack metals other than titanium, and
these processes are electrochemical in nature.
Proponents of this technique argue that implants treated
by sandblasting and acid etch provide superior
radiographic bone densities along implant interfaces
compared with titanium plasma-sprayed surfaces.

Cobalt and Iron Alloys
 The alloys of cobalt (Vitallium) and iron (surgical
stainless steel – 316L) exhibit oxides of chromium
(primarily Cr2O3 with some suboxides) under normal
implant surface finishing conditions after acid or
electrochemical passivation.
 These chromium oxides, as with titanium and alloys,
result in a significant reduction in chemical activity and
environmental ion transfers.
 Under normal conditions of acid passivation, these
chromium oxides are relatively thin (nanometer
dimensions) and have an amorphous atomic structure.
 The oxide atomic spatial arrangement can be
converted to a crystalline order by elevated temperature
or electrochemical exposures.www.indiandentalacademy.comwww.indiandentalacademy.com

 The chromium oxides on cobalt and iron alloys are
microscopically smooth, and again, roughness is usually
introduced by substrate processing (grinding, blasting, or
etching).
 The tissue integration of cobalt alloy could be
described by tissue-to-oxide and tissue-to-metallic
carbide zones. This is uniquely different compared with
titanium implant biomaterials where tissue-to-oxide
regions predominate at the interface.
 The iron-based alloy chromium oxide and substrate
are more susceptible to environmental breakdown, in
comparison to cobalt and titanium-based biomaterials.
However, in the absence of surface damage, the
chromium oxides on stainless steel biomaterials have
shown excellent resistances to breakdown.www.indiandentalacademy.comwww.indiandentalacademy.com

•Dental implants and implant abutments have also been
fabricated from gold alloy with many abutments
fabricated from palladium or Co-Cr-Ni- Mo alloys. The
minimally alloyed gold and palladium systems are noble
electrochemically and do not depend on surface oxides
for chemical and biochemical inertness. However, some
palladium alloys and other lower noble element content
alloys gain chemical and biochemical inertness from
complex metallic surface oxides. As mentioned, the
multicomponent (wrought) cobalt-based alloys, as with
other base-metal systems, depend on chromium oxide
surface conditions for inertness.

In general the noble metal alloys do not demonstrate
the same characteristics of tissue interaction when
compared with the base metal (Ti and Co alloy)
systems.
The ultra structural aspects of tissue integration have
not been extensively investigated for noble alloy
systems. The noble alloys when used in a polished
condition are resistant to debris accumulation on a
relative basis compared with other alloys. This has been
listed as an advantage for their use in intraoral abutment
systems.
Also, mechanical finishing of the more noble alloys
can result in a high degree of polish and a minimal
concern about damaging or removing surface oxides.

Ceramics
Aluminum oxide (Al2O3) ceramics have been
extensively investigated related to surface properties
and how these properties relate to bone and soft tissue
integration.
Bone and soft tissue integration have been
demonstrated for this oxide material over the long term
in humans and laboratory animals. Direct relationships
have been established between the interfacial events of
tissue integration for metallic surface oxides of titanium
and chromium and the aluminum oxide systems.

•Surface quality can be directly correlated with tissue
integration and clinical longevity. Because the aluminum
oxides are crystalline, biomechanical instabilities do not
alter the chemical aspects of biomaterial properties . (No
electrochemical change is introduced if the surface is
removed).
• Ceramic coatings (Al2O3) have been shown to
enhance the corrosion resistance and biocompatibility of
metal implants.
• Studies in orthopedics caution that the Al2O3 coating
may cause a demineralization phenomenon caused by a
high local concentration of substrate ions in the
presence of metabolic bone disease. This remains to be
established within the use of aluminum oxide implants
for clinical applications.

Porous and Featured Coatings
The implant surface may also be covered with a
porous coating. These may be obtained with titanium or
hydroxyapatite particulate – related fabrication
processes.
Titanium Plasma Sprayed
Porous or rough titanium surfaces have been
fabricated by plasma spraying a powder form of molten
droplets at high temperatures. At temperatures in the
order of 15,000C, an argon plasma is associated with a
nozzle to provide very high velocity 600 m/sec partially
molten particles of titanium powder (0.05 to 0.1 mm
diameter) projected onto a metal or alloy substrate. The
plasma sprayed layer after solidification (fusion) is often
provided with a 0.04 to 0.05 mm thickness.www.indiandentalacademy.comwww.indiandentalacademy.com

When examined microscopically, the
coatings show round or irregular pores that
can be connected to each other. These
types of surfaces were first developed by
Hahn and Palich, who reported bone in
growth in plasma spray titanium hybrid
powder plasma spray-coated implants
inserted in animals.
Advantage
increase the total surface area (upto
several times),
produce attachment by osteoformation,
 enhance attachment by increasing ionic
interactions,

introduce a dual physical and chemical anchor system,
and increase the load – bearing capability 25% to 35%.
 The optimum pore size ranged from 150 to 400 µm
and coincidentally correspond to surface feature
dimensions obtained by some plasma spraying
processes.
In addition , porous surfaces can result in an increase
in tensile strength through in growth of bony tissues into
three dimensional features. High shear forces
determined by the torque testing methods and improved
force transfer into the periimplant area have also been
reported.

In 1985 at the Brussels Osseointegration Conference,
the basic science committee did not present results that
showed any major differences between smooth, rough,
or porous surfaces regarding their ability to achieve
osseointegration. However, proponents of porous
surface preparations reported that there have been
results showing faster initial healing compared with
noncoated-porous titanium implants and that porosity
allows bone formation within the porosities even in the
presence of some improvement during the healing
phase. The basic theory was based on increased area
for bone contact.

Reports in the literature caution about cracking and
scaling of coatings because of stresses produced by
elevated temperature processing and risk of
accumulation of abraded material in the interfacial zone
during implanting of titanium plasma sprayed implants.
It may be indicated to restrict the limit of coatings in
lesser bone densities that cause less frictional torque
transfer during implant placement process.

Hydroxyapatite Coating
Hydroxyapatite coating by plasma spraying was
brought to the dental profession by deGroot.
Kay et al. showed with scanning electron microscopy
(SEM) and spectrographic analyses that the plasma-
sprayed HA coating could be crystalline and could offer
chemical and mechanical properties compatible with
dental implant applications.
Block and Thomas showed an accelerated bone
formation and maturation around HA-coated implants in
dogs when compared with noncoated implants.
HA coating can also lower the corrosion rate of the
same substrate alloys.

Cook et al. measured the HA coating thickness after
retrieval from specimens inserted in animals for 32
weeks and showed a consistent thickness of 50 µm,
which is in the range advocated for manufacturing.
The bone adjacent to the implant has been reported to
be better organized than with other implant materials
and with a higher degree of mineralization.
In addition, numerous histologic studies have
documented the greater surface area of bone apposition
to the implant in comparison to uncoated implants, which
may enhance the biomechanics and initial load-bearing
capacity of the system.
HA coating has been credited with enabling HA-coated
Ti or Ti alloy implants to obtain improved bone-to-
implant attachment compared with machined surfaces.www.indiandentalacademy.comwww.indiandentalacademy.com

HA coated
threaded
implant
HA coated
machined
collar cylinder
implant www.indiandentalacademy.comwww.indiandentalacademy.com

Titanium screw
implant with HA
coating

Implants of solid sintered hydroxyapatite have been
shown to be susceptible to fatigue failure. This situation
can be altered by the use of a CPC coating along
metallic substrates. Although several methods may be
used to apply CPC coatings, the majority of
commercially available implant systems are coated by a
plasma spray technique. A powdered crystalline
hydroxyapatite is introduced and melted by a the hot,
high-velocity region of a plasma gun and propelled onto
the metal implant as a partially melted ceramic. One of
the concerns regarding CPC coatings is the strength of
the bond between the CPC and the metallic substrate.
Ion-beam sputtering coating techniques for CPC or
CPC-like nonresorbable coatings to varied substrates
appear to produce dense, more tenacious and thinnerwww.indiandentalacademy.comwww.indiandentalacademy.com

coatings ( a few micrometers), which would minimize the
problem of poor shear strength and fatigue at the
coating-substrate interface. Recent reports have
introduced a new type of treatment for coatings, which
appear primarily amorphous in nature, and further in vivo
studies are needed to determine tissue response. Other
investigations include developing new biocompatible
coatings based on tricalcium phosphate or titanium
nitride.
It has been shown that the plasma-spraying
technique can alter the nature of the crystalline ceramic
powder and can result in the deposition of a variable
percentage of a resorbable amorphous phase. A dense
coating with a high crystallinity has been listed as
desirable to minimize in vivo resorption.www.indiandentalacademy.comwww.indiandentalacademy.com

In addition, the deposited CPC may be partially resorbed
through remodeling of the osseous interface. It is
therefore wise to provide a biomechanically sound
substructure design that is able to function under load-
bearing conditions to compensate for the potential loss
of the CPC coating over years. In addition, the CPC
coatings may resorb in infected or chronic inflammation
areas.
One advantage of CPC coatings is that they can act as a
protective shield to reduce potential slow ion release
from the Ti-6Al-4V substrate. Also, the interdiffusion
between titanium and calcium, and phosphorus and
other elements may enhance the coating substrate bond
by adding a chemical component to the mechanical
bond. www.indiandentalacademy.comwww.indiandentalacademy.com

Advantage of TPS and HA
 Increased surface area
 Increased roughness for initial stability
 Stronger bone to implant interface
 Faster healing bone interface
 Increased gap healing between bone and HA
 Stronger interface than TPS
 Less corrosion of metal
Disadvantage
• Flaking, cracking or scaling upon insertion
• Increased plaque retention
• Increased bacteria and nidus for infection
• Complication of treatment of failing implants
• Increased cost www.indiandentalacademy.comwww.indiandentalacademy.com

Other Surface Modifications
Surface modification methods include controlled
chemical reactions with nitrogen or other elements or
surface ion implantation procedures. The reaction of
nitrogen with titanium alloys at elevated temperatures
results in titanium nitride compounds being formed along
the surface. These nitride surface compounds are
biochemically inert (like oxides) and alter the surface
mechanical properties to increase hardness and
abrasion resistance. Most titanium nitride surfaces are
gold in color, and this process has been extensively
used for enhancing the surface properties of industrial
and surgical instruments.

Increased hardness, abrasion, and wear resistance can
also be provided by ion implantation of metallic
substrates. The element most commonly used for
surface ion implantation is nitrogen. Electrochemically,
the titanium nitrides are similar to the oxides (TiO2),and
no adverse electrochemical behavior has been noted if
the nitride is lost regionally. The titanium substrate
reoxidzes when the surface layer of nitride is removed.
Nitrogen implantation and carbon-doped layer deposition
have been recommended to improve the physical
properties of stainless steel without affecting its
biocompatibility. Again, questions could be raised about
coating loss and crevice corrosion.

Surface Cleanliness
A clean surface is an atomically clean surface with
no other elements than the biomaterial constituents.
Contaminants can be particulates, continuous films (oil,
fingerprints), and atomic impurities or molecular layers
(inevitable) caused by the thermodynamic instability of
surfaces. Even after reacting with the environment,
surfaces have a tendency to lower their energy by
binding elements and molecules. The typical
composition of a contaminated layer depends on
atmospheres and properties of surface. For example,
high-energy surfaces ( metals, oxides, ceramics) usually
tend to bind more to this type of monolayer than
polymers and carbon (amorphous).

In the earlier times of dental implantology, no specific
protocol for surface preparation, cleaning, sterilization,
and handling of the implants we established.
Baier et al. and Kasemo et al. have respectively
demonstrated adverse host responses caused by faulty
preparation and sterilization, omiation to eliminate
adsorbed gases, and organic and inorganic debris.
 According to Albrektsson et al., implants that seem
functional may fail even after years of function and the
cause may be attributed to improper ultrasonic cleaning,
sterilization, or handling during the surgical placement.
A systematic study of contamination layers is not
available. Lausnaa et al. showed that titanium implants
had large variations in carbon contamination loads (20%
to 60%) in the 0.3 to 1 nm thickness rangewww.indiandentalacademy.comwww.indiandentalacademy.com

attributed to air exposure and residues from cleaning
solvents and lubricants used during fabrication. Trace
amounts of Ca,P,N,Si,S,C1, and Na were noted from
other studies. Residues of fluorine could be attribted to
passivation and etching treatments; Ca, Na, and C1 to
autoclaving; and Si to sand and glass beading
processes

Surface Energy
Measurements of surface property values of an
implant's ability to integrate within bone include contact
angle with fluids, local pH, and surface topography.
These are often used for the determination of surface
characteristics.
Baier et al. conducted numerous studies to evaluate
liquid, solid, and air contact angles, wetting properties,
and surface tensions as criteria to assess surface
cleanliness because these parameters have been
shown to have direct consequence on osseointegration.
High surface energy is said to be most desirable. High
surface energy implants showed a threefold increase in
fibroblast adhesion and higher energy surfaces such as
metals, alloys, and ceramics are best suited to achieve
cell adhesion. www.indiandentalacademy.comwww.indiandentalacademy.com

Surface tension values of 40 dyne/cm and higher are
characteristic of very clean surfaces and excellent
biologic integration conditions. A shift in contact angle
(increase) is related to the contamination of the surface
by hydrophobic contaminants and decreases the surface
tension parameters. Because a spontaneously
deposited, host-dependent “conditioning film” is a
prerequisite to the adhesion of any biologic element, it is
suggested that the wetting of the surface by blood at the
time of placement can be a good indication of the high
surface energy of the implant.

Passivation and Chemical Cleaning
The ASTM (ASTM B600, ASTM F-86)
specifications for final surface treatment of surgical
titanium implants require pickling and descaling with
molten alkaline base salts. This is often followed by
treatment with a solution of nitric or hydrofluoric acid to
decrease and eliminate contaminants such as iron. Iron
or other elements may contaminate the implant surface
as a result of the machining process. This type of debris
can have an effect of demineralizing of the bone matrix.
But these finishing requirements remain very general.
Studies of fibroblast attachment on implant surfaces
showed great variations depending on the different
processes of surface preparation. Inoue et al. showed
fibroblasts developed a capsule or oriented fibrous
attachment following the grooves in titanium disks.www.indiandentalacademy.comwww.indiandentalacademy.com

Contact angles are also greatly modified by acid
treatment or water rinsing. Machining operations,
polishing, texturing process, residual chemical deposits,
and alloy microstructure all inadvertently affect the
surface composition. There are also many ways to
intentionally modify the surface of the implant. They
include conventional mechanical treatment (sand
blasting), wet or gas chemical reaction treatment,
electroplating or vapor plating, and ion-beam
processing, which leaves bulk properties intact and has
been newly adapted to dentistry from thin film
technology. A general rule has been that cleaner is
better.

Sterilization
Manipulation with bare fingers or powdered
gloves, tap water, and residual vapor-carried debris from
autoclaving can all contaminate implant surfaces.
Bauhammers, in an SEM study of dental implants,
showed contamination of the surface with acrylic
materials, powder for latex gloves, and bacteria.
Today, in most cases, the manufacturer guarantees
precleaned and presterilized implants with high
technology procedures, with the implants ready to be
inserted. If an implant needs to be resterilized,
conventional sterilization techniques are not normally
satisfactory. It appears at the present time that no
sterilization medium is totally satisfactory for all
biomaterials and designs.www.indiandentalacademy.comwww.indiandentalacademy.com

Baier et al. showed that steam sterilization can cause
deposits of organic substances resulting in poor tissue
adhesion.
Doundoulakis submitted Ti samples to different
sterilization techniques, concluded to the adverse effect
of steam sterilization and degradative effect of
endodontic glass bead sterilizers, found that dry heat
sterilization leaves organic deposits on the surface and
suggested that UV light sterilization may become a good
alternative after further evaluation.

In addition, accelerated oxide growth on Ti may occur
with impurity contamination leading to surface
discoloration.
In a study by Draughn et al., corrosion products and
films from autoclaving, chemicals, and cytotoxic residues
from solutions were identified at the surface of implants
submitted to sterilization. They suggested that alteration
of the Ti surface by sterilization methods may in turn
affect the host response and adhesive properties of the
implant.
On the other hand, Schneider et al. compared the
surface of Ti plasma-sprayed and HA-coated Ti implants
after steam or ethylene dioxide sterilization using
energy depressive x-ray analysis and concluded that
these techniques do not modify the elemental
composition of the surface.www.indiandentalacademy.comwww.indiandentalacademy.com

Keller et al. studied the growth of fibroblasts on disks of
CP titanium sterilized by autoclaving, ethvleneoxide,
ethyl alcohol, or solely passivated with 30% nitric acid
and concluded that sterilization seems to inhibit cell
growth, whereas passivation does not.
Presently, proteinaceous deposits can be best
eliminated by radio-frequency glow discharge technique
(RFGDT), which seems to be a suitable final cleaning
procedure. The implants are treated within a controlled
noble gas discharge at very low pressure. The gas ions
bombard the surface and remove surface atoms and
molecules, which are absorbed onto it or are
constituents of it. However, the quality of the surface
treated depends on the gas purity.

Baier et al. showed that RFGDT is good for cleaning and
at the same time, for granting a high energy state to the
implant, which is related to improved cell adhesion
capabilities. Thinner, more stable oxide films and
cleaner surfaces have been reported with RFGDT plus
improved wet ability and tissue adhesion. The principal
oxide at the surface is unchanged by the RFGDT
process. A decrease in bacteria contamination of HA-
coated implant surfaces was reported after RFGDT, and
studies suggest that RFGDT may enhance calcium
and/or phosphate affinity because of an increase in
elemental zone at the surface resulting in the formation
of amorphous calcium phosphate compounds

Lately, a modified ultraviolet (UV) light sterilization
protocol showed to enhance bioreactivity, which was
also effective for eliminating some biological
contaminants.
Singh and Schaaf assessed the quality of UV light
sterilization and its effects on irregularly shaped objects,
and they established it s effectiveness on spores and its
ability to safely and rapidly clean the surface and to
grant high surface energy.
Hartmand et al. submitted implants to various
pretreatment protocols (RFGDT, UV light, or steam
sterilization) and inserted them in miniature swine.
Although RFGDT and UV-sterilized implants showed
rapid bone ingrowths and maturation, steam sterilized
implants seemed to favor thick collagen fibers at the
surface. www.indiandentalacademy.comwww.indiandentalacademy.com

On the other hand, Carlsson et al. inserted implants in
rabbits and compared the performances of
conventionally treated implants with implants treated
with RFGDT, found similar healing responses, and
further cautioned that the RFGDT process produces a
much thinner oxide layer at the surface of the implant
and may deposit silica oxide from the glass envelope.
Adequate sterilization of clean, prepackaged
dental implants and related surgical components has
resulted in an ever expanding use of gamma radiation
procedures. Because gamma radiation sterilization of
surgical implants is a well-established methodology
within the industry, facilities, procedures, and standards
are well known.

Most metallic systems are exposed to radiation doses
exceeding 2.5 megarads where the packaging and all
internal parts of the assembly are sterilized. This is an
advantage in that components remain protected, clean,
and sterile until the inner containers are opened within
the sterile field of the surgical procedure. The healing
screws, transfer elements, wrenches, and implants are
all exposed to the gamma sterilization, which reduces
opportunities for contamination.
Some ceramics can be discolored and some
polymers degraded by gamma radiation exposures. The
limits are known for classes of biomaterials and all types
of biomaterials ca be adequately sterilized within the
industry. Systems control, including prepackaging and
sterilization, has been an important part of the success
of dental implantology.www.indiandentalacademy.comwww.indiandentalacademy.com

SUMMARY
The biomaterials discipline has evolved
significantly over the past decades, and synthetic
biomaterials are now constituted, fabricated, and
provided to health care professionals as mechanically
and chemically clean devices that have a high
predictability of success when used appropriately within
the surgical disciplines.
Surface characterization and working knowledge
about how surface and bulk biomaterial properties
interrelate to dental implant biocompatibility profiles
represent an important area in implant-based
reconstructive surgery.

REVIEW OF LITERATURE
Bothe et al (1940) tested a wide spectrum of metals for orthopedic
use and found that titanium has a special relation with bone.
Branemark (1960) suggested the possibility of osseointegration
using titanium implants.
Reisbick and Benson (1974) conducted a study on ceramic coated
subperiosteal implants and noticed direct attachment to alumina
surface.

Rostoker et al.(1974) studied couple corrosion in vitro for alloys
and found that dissimilar metals in a combined prosthesis did not
create a regional breakdown of the titanium passive layer. A
second in vivo study evaluated couple/crevice corrosion of
prosthetic alloys in vertebral muscles of dogs for 30 weeks (non –
load bearing, nonosseointegrated). It was concluded that metals
of superior corrosion resistance, such as titanium alloy, and
wrought cobalt alloys can be combined with titanium alloy in one
prosthesis to provide superior mechanical performance without
creating additional corrosion. However, repeated oxide breakdown
such as sustained abrasion was likely to damage the corrosion
resistance of an alloy for any type of coupling.

In 1981 Clemow et al. showed that the rate and percentage of
bone ingrowth into the surface of TPS was inversely proportional
to the square root of the pore size for sizes greater than 100 µm
and that the shear properties of the interface were proportional to
the extent of bone ingrowth.
Parr et al (1985) stated that titanium and its alloys posses
mechanical properties that make them ideal for implant materials.
Titanium oxodises in air and the stability and inertness of this
surface oxide layer acts to protect titanium from corrosive
biocorrosion.
Geis Gestorfer(1988) used linear polarization methods to show
that titanium showed minimal breakdown in simulated tissue fluids
whereas Ni-Ti showed rapid breakdown of passivity with
increased chlorine product related concentrations in unbuffered
solutions.Therefore body fluids could be responsible for the
dissolution of some metallic passive oxide films.www.indiandentalacademy.comwww.indiandentalacademy.com

Zetterquist at al (1995) studied the interface between ceramic
implant and bone and concluded that oeesointegration can be
obtained. And no leakage of aluminum could be detected in bone.
Denissen et al (1996) studied Calcium phosphate ceramic
coatings with a hydroxyapatite chemistry applied on the surface of
dental implants to eliminate the need for initial mechanical
retention and decrease the time necessary for bonding the
implants to the bone. Hydroxyapatite-coated implants retrieved
from patients were found to be compatible and to have bonded
strongly to the bone, but the coatings showed thinning because of
partial or total loss of coating material.

Cross-Poline GN et al (1997) compared the surface roughness
produced by various implant curets on titanium implant abutment
surfaces. The three experimental surfaces were instrumented with a
gold platinum curet, an unreinforced resin curet, or a reinforced resin
curet. Two implants were assigned to each of the following treatments:
128, 256 or 512 scaling strokes within a 4 mm wide area. Photographs
were taken of the surfaces with a scanning electron microscope The
surfaces were different at 8 and 16 years. At 8 years, the surface
roughness was significant between the treatments in the following
ascending order: untreated, unreinforced resin curet, reinforced resin
curet and gold platinum curet. Significant roughness was observed for
surfaces treated by only the gold platinum curet and the reinforced resin
curet at 16 years. The gold platinum curet created the roughest surface.

P.X. Holding et al (1998) stated that Fluoride ions are the only
aggressive ions for the protective oxide layer of titanium and
titanium alloys. Thus their presence may possibly start a localized
corrosive degradation by pitting and crevice corrosion processes.
Since hygiene products like toothpastes and prophylactic gels
contain fluoride ions, Two different milieu based on the Fusayama
artificial saliva and an electrolyte solution containing NaCl, with
and without fluoride ions, was used. Results showed (a) with and
without fluoride ions, galvanic currents are weak (b) titanium
submitted to anodic polarization in an electrolyte, even one
containing fluoride, merely develops an oxide layer and does not
corrode within that same pH range (c) in confined areas where
fluoride ions are present, titanium and the dental alloys tested
undergo as corrosive process, in the form of crevice and pitting,
as soon as the pH drops below 3.5.

Sawase et al (2001) conducted a study was to examine the
effectiveness of a thick oxide layer on corrosion resistance in vitro and
the bone formation around titanium implants in vivo. A plasma source
ion implantation (PSII) method was used to increase the thickness of the
surface oxide layer. The results indicate that in spite of improved
corrosion resistance in vitro, a thick oxide layer fabricated with the PSII
method does not influence early bone formation around titanium
implants in vivo.

Fathi et al (2003) evaluated the corrosion behavior and thus the
biocompatibility of the uncoated and coated stainless steels and
compared the effect of type of coatings on corrosion behavior. They
used Three types of coatings, hydroxyapatite (HA), titanium (Ti), and a
double-layer HA/Ti on 316L stainless steel. HA coating was produced
using plasma-spraying technique and Ti coating was made using
physical vapor deposition process. In order to perform a novel double-
layer composite coating, a top layer of HA was plasma-sprayed over a
physical vapor deposited Ti layer on 316L stainless steel. Results
showed that Double-layer HA/Ti coating on 316L SS had a positive
effect on improvement of corrosion behavior. The decrease in corrosion
current densities was significant for these coated specimens and was
much lower than the values obtained for uncoated and single HA coated
specimens. Ti coating on 316L SS also has a beneficial effect on
corrosion behavior. These results demonstrated that the double-layer
HA/Ti coated 316L SS can be used as an endodontic implant and two
goals including improvement of corrosion resistance and bone
osteointegration can be obtained simultaneously.

Nogueras-Bayona et al (2004) examined the roughness and
bonding strength of the chemical-made apatite layer in comparison with
the titanium surface and the plasma-sprayed apatite. Commercially pure
titanium plates were heated and chemically treated to deposit crystalline
apatite on their surface. The roughness of the titanium surface of the
original samples and the apatite surface was analyzed by a roughness
surface tester. A scratch test was used to compare the adhesion of the
chemical apatite layer to the titanium with the adhesion of a plasma-
sprayed layer. A dense bone-like apatite layer was formed on the
surface of the titanium by a simple chemical method. The surface
roughness test showed that the chemical apatite coating increased the
roughness of the samples. The scratch test showed that the bonding
strength of the chemical-made apatite coatings to the titanium substrate
was higher than the plasma-sprayed apatite coatings. The apatite layer
produced by chemical treatment did not show a lower roughness than
the titanium substrate. This chemical apatite layer also bonded tighter to
the titanium than the plasma-sprayed apatite. This chemically made
apatite coating is expected to provide a long-term implant-bone fixation.

 The Dental Implant- Clinical And Biological Response Of Oral
Tissues. Ralph McKinney and Jack Lemons.
 Vitreous carbon implants- Paul Schnitman and Leonard
Schulman. DCNA July 1980.
 Pyrolytic carbon and carbon coated metallic dental implants-
JohnKent and Jack Bokros. DCNA July 1980.
 Miller RJ. Treatment of the contaminated implant surface using
the Er,Cr:YSGG laser. Implant Dent. 2004 Jun;13(2):165-70.
 Shirakura M, Fujii N, Ohnishi H, Taguchi Y, Ohshima H,
Nomura S, Maeda T. Tissue response to titanium implantation in
the rat maxilla, with special reference to the effects of surface
conditions on bone formation. Clin Oral Implants Res. 2003
Dec;14(6):687-96.
 Gluszek J, Jedrkowiak J, Markowski J, Masalski J.Galvanic
couples of 316L steel with Ti and ion plated Ti and TiN coatings in
Ringer's solutions. Biomaterials. 1990 Jul;11(5):330-5.

 Reclaru L, Meyer JM. Effects of fluorides on titanium and other
dental alloys in dentistry. Biomaterials. 1998 Jan-Feb;19(1-3):85-
92.
 Nogueras-Bayona J, Gil FJ, Salsench J, Martinez-Gomis
Roughness and bonding strength of bioactive apatite layer on
dental implants. Implant Dent. 2004 Jun;13(2):185-9.
 Augthun M, Tinschert J, Huber A. In vitro studies on the effect
of cleaning methods on different implant surfaces. J Periodontol.
1998 Aug;69(8):857-64.
 Cross-Poline GN, Shaklee RL, Stach DJ.Effect of implant curets
on titanium implant surfaces. Am J Dent. 1997 Feb;10(1):41-5.
 Catledge SA, Fries MD, Vohra YK, Lacefield WR, Lemons JE,
Woodard S, Venugopalan R. Nanostructured ceramics for
biomedical implants. J Nanosci Nanotechnol. 2002 Jun-Aug;2(3-
4):293-312.
 Denissen HW, Klein CP, Visch LL, van den Hooff A. Behavior of
calcium phosphate coatings with different chemistries in bone. Int J
Prosthodont. 1996 Mar-Apr;9(2):142-8.www.indiandentalacademy.comwww.indiandentalacademy.com

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