2. 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.
3. 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.
4. 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.
5. 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
6. 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.
7. •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.
8. •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.
9. 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.
10. 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.
11. 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
12. 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
13. 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
14. 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.
15. 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.
16. 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.
17. 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.
20. 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.
21. 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.
22. 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.
23. 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.
24. 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.
25. 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.
27. 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.
28. 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.
29. 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.
31. 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.
32. 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.
33. 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.
34. 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.
35. 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.
36. 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.
37.
38. 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 silicon
• Vitreous and Pyrolytic
40. 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.
41. 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.
42. 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).
43. 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.
44. 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.
45. 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.
46. 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.
47. 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.
48. 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.
49. 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.
51. 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.
52. 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.
54. 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.
55. 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).
56. 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.
57. 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.
58. 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).
59.
60. 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.
61. 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. (The latter has been carefully
cleaned in vacuum before analysis). A large carbon
signal (- 40 atomic %) is always observed, as well as a
smaller nitrogen one. 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.
62. 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, but at least the C signal
may originate from surface segregation of a low
concentration of Ca in the titanium sample.
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.
63. 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.
64. 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.
65. 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.
66. 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.
67. 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.
68. 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.
69. 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; while molybdenum
provides strength and bulk corrosion resistance.
70. 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.
71. 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 many of these alloy devices have
been fabricated by dental laboratories, all aspects of
quality control and analysis for surgical implants must be
followed during allow selection, casting, and finishing.
Critical considerations include the chemical analysis,
mechanical properties, and surface finish as specified by
the ASTM Committee F4 on surgical implants and the
ADA. When properly fabricated, implants from this alloy
group have shown excellent biocompatibility profiles.
72. Because of the requirements of low cost and long term
clinical success these alloys have been used
extensively in many areas of surgery and dentistry.
However the greater corrosion resistance and tissue
compatibility of titanium have made it a particularly
effective metal for dental implants.
73. 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.
74. 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 minimized 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. In some clinical conditions,
more than one alloy may be present within the same
dental arch of a patient.
75. 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.
76. 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. Some significant
advantages of these reactive group metals and their
alloys have been reported, although large numbers of
such devices have not been fabricated in the United
States.
Gold, platinum, and palladium are metals of relatively
low strength, which places limits on implant design.
77. These metals, especially gold because of nobility and
availability, continue to be used as surgical implant
materials.
78. CERAMICS AND CARBON
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.
79. Aluminum, Titanium And Zirconium Oxides
High 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.
80. 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.
82. 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.
83.
84. Bioactive and Biodegradable Ceramics Based on
Calcium Phosphates
Bone Augmentation and Replacement
The calcium phosphate (CaPO4) ceramics used in
dental reconstructive surgery include a wide range of
implant types and thereby a wide range of clinical
applications. Early investigations emphasized solid and
porous particulates with nominal compositions that are
relatively similar to the mineral phase of bone
(Ca5[PO4]3OH). Microstructural and chemical
properties of these particulates were controlled to
provide form that would remain intact for structural
purposes after implantation.
85. The laboratory and clinical results for these particulates
were most promising and led to expansions for implant
applications, including larger implant shapes (such as
rods, cones, blocks, H-bars) for structural support under
relatively high-magnitude loading conditions. Also, the
particulate size range for bone replacements was
expanded to both smaller and larger sizes for combined
applications with organic compounds. Mixtures of
particulates with collagen, and subsequently with drugs
and active organic compounds such as bone
morphogenetic protein (BMP), increased the range of
possible applications. Over the past 20 years, these
types of products and their uses have continued to
significantly expand.
86. Endosteal and Subperiosteal Implants
The first series of structural forms for dental implants
included rods and cones for filling tooth root extraction sites
(ridge retainers) and, in some cases, load-bearing endosteal
implants. Limitations in mechanical property characteristics
soon resulted in internal reinforcement of the CaPO4
ceramics implants through mechanical (central metallic
rods) or physicochemical (coating over another substrate)
techniques.
The coatings of metallic surfaces using flame or
plasma spraying (or other techniques) increased rapidly for
the CaPO4 ceramics. The coatings have been applied to a
wide range of endosteal and subperiosteal dental implant
designs with an overall intent of improving implant surface
biocompatibility profiles and implant longevities.
87. Advantages
Chemical compositions of high purity and of
substances that are similar to constituents of normal
biologic tissue (calcium, phosphorus, oxygen, and
hydrogen)
Excellent biocompatibility profiles within a variety of
tissues, when used as intended
Opportunities to provide attachments between
selected CaPO4 ceramics and hard and soft tissues
Minimal thermal and electrical conductivity plus
capabilities to provide a physical and chemical barrier to
ion transport (e.g., metallic ions)
Moduli of elasticity more similar to bone than many
other implant materials used for load-bearing implants
88. • Color similar to bone, dentin, and enamel
• As evolving and extensive base of information related
to science, technology, and application.
Disadvantages
Variations in chemical and structural characteristics for
some currently available implant products
Relatively low mechanical tensile and shear strengths
under condition of fatigue loading.
Relatively low attachment strengths for some coating-
to-substrate interfaces.
Variable solubility’s depending on the product and the
clinical application. The structural and mechanical
stabilities of coatings under in vivo load-bearing
conditions (especially tension and shear may be
variable as a function of the quality of the coating.
89. Alterations of substrate chemical and structural
properties related to some available coating technologies
Expansion of applications that sometimes exceed the
evolving scientific information on properties.
In general, these classes of bioceramics have
lower strengths, hardness, and moduli of elasticity than
the more chemically inert forms previously discussed.
Fatigue strengths, especially for porous materials, have
imposed limitations with regard to some dental implant
designs. In certain instances, these, characteristics have
been used to provide improved implant conditions (e.g.,
biodegradation of particulates). Calcium aluminates,
sodium-lithium invert glasses with calcium phosphate
additions (Bioglass or Ceravital, and glass ceramics (AW
glass-ceramic) also provide a wide range of properties
and have found extended applications.
90.
91. One of the more important aspects of the CaPO4
ceramics relates to the possible reactions with water. For
example, hydration can convert other compositions to
HA: also, phase transitions among the various structural
forms can exist with any exposure to water. This has
caused some confusion in the literature, in that some
CaPO4 ceramics have been steam autoclaved for
sterilization purposes before surgical implantation.
Steam or water autoclaving can significantly change the
basic structure and properties of CaPO4 ceramics ( or
any bioactive surface) and thereby provide an unknown
biomaterial condition at the time of implantation. This is
to be avoided through the use of presterilized or clean,
dry heat or gamma sterilized conditions.
92. The two calcium phosphate systems that have been
most investigated as bone implant materials are HA and
TCP. Based on numerous experiments it was apparent
that HA ceramics could be considered to be long term or
permanent bone implant materials, whereas porous TCP
ceramics could serve as bioresorbables.
Forms, Microstructures, and Mechanical
Properties
Particulate HA, provided in a nonporous (<5%
porosity) form as angular or spherically shaped
particles, is an example of a crystalline, high-purity HA
biomaterial These particles can have relatively high
compressive strengths (up to 500 Mpa), with tensile
strengths in the range of 50 to 70 Mpa.
93. Usually, dense polycrystalline ceramics consisting of
small crystallites exhibit the highest mechanical
strength, apart from monocrystalline ceramics free of
defects (such as single crystal sapphire implants).
Ceramics are brittle materials and exhibit high
compressive compared with tensile strengths.
However, less resistance to tensile and shear stresses
limit their application as dental implants because of
mechanical constraints of implant form and volume.
Nonresorbable, “bioinert” ceramics exhibiting
satisfactory load-bearing capability are limited to dense
monocrystalline and poly-crystalline aluminum, irconium,
and titanium oxide ceramics. These same mechanical
characteristics exist for the solid portions of several
porous HA particulates and blocks.
94. The porous materials also provide additional regions for
tissue ingrowths and integration (mechanical
stabilization) and thereby a minimization of interfacial
motion and dynamic (wear-associated) interfacial
breakdown. The strength characteristics after tissue in
growth would then become a combination of the
ceramic and the investing tissues.
A number of the CaPO4 ceramics are phase
mixtures of HA and TCP, whereas some compounds are
composites or mechanical mixtures with other materials.
These classes of bioactive ceramics, including glasses,
glass-ceramics, mixtures of ceramics, combinations of
metals and ceramics, or polymers and ceramics, exhibit
a wide range of properties.
95. In general, these biomaterials have shown acceptable
biocompatibility profiles from laboratory and clinical
investigations. Bulk-form implant designs made from
calcium phosphate ceramics, which were shown to be
contraindicated for some implant designs because of
poor mechanical performance, have found a wide range
of indications as coatings of stronger implant materials.
The coatings of CaPO4 ceramics onto metallic
(Co- and Ti-based) biomaterials have become a routine
application for dental implants. These coatings for the
most part are applied by plasma spraying, have average
thickness between 50 and 70 um, are mixtures of
crystalline and amorphous phases, and have variable
microstructures (phases and porosities) compared with
the solid portions of the particulate forms of HA and TCP
biomaterials.
96. Concerns continue to exist about the fatigue strengths of
the CaPO4 coatings and coating-to-substrate interfaces
under tensile and shear loading conditions. There have
been some reports of coating loss as a result of
mechanical fracture, although the numbers reported
remain small. This has caused some clinicians and
manufacturers to introduce designs in which the coating
are applied to shapes (geometric designs) that minimize
implant interface shear or tensile loading conditions (
such as porosities, screws, spirals, plateaus, and vents).
From theoretic considerations, the coating of
mechanically protected areas seems most desirable.
97. Density, Conductivity, and Solubility
Bioactive ceramics are especially interesting for
implant dentistry because the inorganic portion of the
recipient bone is more likely to grow next to a more
chemically similar material. Under the bioactive
(bioreactive) categorization are included calcium
phosphate materials such as TCP, HA, calcium carbonate
(corals), and calcium sulfate-type compounds and
ceramics. A chemical-biochemical, contact between the
host bone and grafted material may be developed as
well as a possible stimulus of bone activity. Their
limitations have been associated with the material forms
that have lower strengths (i.e., similar to or less than
bone).
98. Dissolution characteristics of bioactive ceramics have
been determined for both particulates and coatings. In
general, solubility is greater for TCP than for HA. Each
increase relative to increasing surface area per unit
volume (porosity) and the CaPO4 ceramic solubility
profiles depend on the environment (pH, mechanical
motion, etc.).
If one considers a uniform material chemistry, the larger
the particle size, the longer the material will remain at
an augmentation site. Thus 75 um size particles will be
resorbed more rapidly than 3000 um size particles. Also,
the porosity of the product impacts the resorption rate.
99. Tofe et al. reported on the porosity of dense,
macroporous and microporous calcium phosphates.
Some of the dense HA lacks any macro- or microporosity
within the particles. The longest resorption rate
occurred with the dense nonporous HA type because
osteoclasts may only attack the surface and cannot
penetrate the nonporous material. Macroporous
calcium phosphates ( such as corallin HA) demonstrated
100 um or 500 um pores, which composed 15% or more
of the total material volume.
100. Minimal porosity was found in the HA bulk material that
surrounded the large pores. Microporous apatites often
have their origin from bovine or human bone. The
porosity observed in these materials is approximately 5
um or less and composes less than 28% of the total
volume. The pores or holes are regions where blood
components and organic materials can reside when
placed within bone and represent the regions where
living material existed before the explanation and
processing of the implant material. The greater the
porosity, the more rapid the resorption of the graft
material.
101. The crystallinity of HA also affects the resorption rate
of the material. The highly crystalline structure is more
resistant to alteration and resorption. An amorphous
product has a chemical structure that is less organized
with regard to atomic structure. The hard or soft
tissues of the body are more able to degrade the
components and resorb the amorphous forms of
grafting materials. Thus crystalline forms of HA are
found to be very stable over the long terms under
normal conditions, whereas the amorphous structures
are more likely to exhibit resorption and susceptibility
to enzyme- or cell-mediated breakdown. Therefore in
general, the less crystalline the material, the faster its
resorption rate. The resorption of the bone substitute
may be cell or solution mediated.
102. Cell-mediated resorption required processes associated
with living cells to resorb the material, similar to the
modeling/remodeling process of living bone, which
demonstrates the coupled resorption/formation process.
A solution mediated resorption permits the dissolution of
the material by a chemical process. Impurities or other
compounds in bioactive ceramcis, such as calcium
carbonate, permit more rapid solution mediated
resorption, which then increases the porosity of the
bone substitute. Hence although the coralline HA does
not demonstrate micropores along the larger holes, the
HA may have carbonates incorporated within the
material, which hastens the resorption process.
103. The pH in the region in which the bone substitutes are
placed also affects the rate of resorption. As the pH
decreases (for example, because of chronic inflammation
or infection) the components of living bone, primarily
the calcium phosphates, resorb by a solution-mediated
process (i.e., they become unstable chemically),
The CaPO4 coatings are nonconductors of heat
and electricity. This can provide a relative benefit for
coated dental implants where mixtures of conductive
materials may be included in the overall prosthetic
reconstruction. In combination with color (off-white),
these properties are considered to be advantageous.
104. Current Status and Developing Trends
The CaPO4 ceramics have roved to be one of the
more successful high technology-based biomaterials that
has evolved within the past decades. Their
advantageous properties strongly support the expanding
clinical applications and the enhancement of the
biocompatibility profiles for surgical implant uses.
Within the overall theme for new generation
biomaterials to be chemically (bonding to tissue) and
mechanically (nonuniform, multidirectional properties)
anisotropic, the CaPO4 ceramics could be the biomaterial
surfaces of choice for many device applications.
105. 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.
106. Ceramic and carbonitic substances continue to be used
as coatings on metallic and ceramic materials.
Advantages are tissue attachment; components that are
normal to physiological environments; 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. Possible 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.
107. 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.
109. Pyrolytic carbon implants
Since vitreous carbon is a brittle material with limited
strenght, 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 strenght 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.
110.
111. Therefore carbon implant can bend and displace as if it
were cortical bone, thus minimising 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.
112. POLYMERS AND COMPOSITES
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 the subatomic scale. These
characteristics are critical to the surface composition,
corrosion resistance, cleanliness, surface energy, flexure,
and tendency to interact, such as the ability to denature
proteins.
113. SURFACE CHARACTERIZATION AND TISSUE
INTERATION
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.
114. 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. 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).
115. 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. 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).
116. 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.
117. 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.
118. 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.
119. The chromium oxides on cobalt and iron alloys are
microscopically smooth, and again, roughness is usually
introduced by substrate processing (grinding, blasting,
or etching). Because these oxides, similar to titanium
oxides, are very thin (nanometer dimensions), the
reflected light color of the alloys depends on the metallic
substrate under the oxide. However, as mentioned, the
titanium, cobalt, and iron metallic systems depend on
the surface reaction zones with oxygen (oxides) for
chemical and biochemical interness.
120. The cobalt and iron alloy bulk microstructures are
normally mixtures of the primary alloy phases with
regions of metallic carbides distributed throughout the
material. Along the surfaces, the chromium oxide covers
the matrix phase (metallic regions). While the carbides
stand as secondary components (usually as mounds
above the surface) at the microscopic level. In contrast
to homogenization annealed alloys, the as-cast cobalt
alloys exhibit mutiphasic characteristics within their
microstructure, with relatively extensive regions of the
alloy surfaces occupied by complex metallic carbides.
Thus 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.
121. The iron-based alloy chromium oxide and substrate
are more susceptible to environmental breakdown, in
comparison to cobalt and titanium-based biomaterials.
In general, if stainless steel implant surfaces are
mechanically altered during implantation, or if the
construct introduces an interface that is subjected to
biomechanical fretting, the iron alloy will biodegrade in
vivo, and the fatigue strength of surgical stainless steel
can be significantly decreased in a corrosive
environment. 180 In some cases this has resulted in
implant loss. However, in the absence of surface
damage, the chromium oxides on stainless steel
biomaterials have shown excellent resistances to
breakdown, and multiple examples of tissue and host
biocompatibility have been shown for implants removed
after long-term (beyond 30 years in vivo) implantations.
122. 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. This would be
the case for the high-noble alloys (major compositions of
gold, platinum, palladium, iridium, and ruthenium).
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.
123. 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, although some bare presented
results describing osseointegration of gold alloys. 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.
124. Ceramics
As mentioned previously , surface quality can be directly
correlated with tissue integration and clinical longevity.
Because the aluminum oxides are crystalline and extend
throughout the surface and bulk zones, 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, in
particular surgical stainless steel and Ni-Cr,Co-Cr alloys.
125. However, the Ni-Cr and steel alloys can be subject to
crevice corrosion. However, 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.
126. Hydroxyapatite
In addition to the bulk aluminum oxide
biomaterials, calcium phosphate –based ceramic or
ceramic-like coatings have been added to titanium and
cobalt alloy substrates to enhance tissue integration and
biocompatibility. These coatings, for the most part, are
applied by plasma spraying small size particles of
crystalline hydroxyapatite ceramic powders.
Surface roughening by particulate blasting can be
achieved by different media. Sandblasting provides
irregular rough surfacing with <10m scales and a
potential for impurity inclusions.
128. Niznick used a titanium alloy Ti-6Al-4V to improve the
mechanical properties and elected to eletropolish the
surface to reduce surface roughness to be only in the
0.1 m scale by controlled removal of the surface layer
by dissolution. 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.
129. Recently, concerns have been expressed regarding
embedded media from glass beading(satin finish) and
grit blasting (alumina Al2O3) and a possible risk of
associated osteolysis caused by foreign debris. Ricci
reported on failed retrieved implants that exhibited
extensive surface inclusions consisting of silicon and/or
aluminium oxide related product, which were alos
present in the surrounding tissues. A relatively new
process (resorbable blast media) has been said to
provide a comparable roughness to an alumina grit blast
finish, which can be a rougher surface than the
machined ,glass beaded, or acid etched surfaces.
130. 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 15,0000C, 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.
131. 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. A porous titanium
surface from various fabrication methods
may 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%.
132. The optimum pore size was deduced from the maximum
fixation strength measurements. These surface
porosities ranged from 150 to 400 m 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.
133. 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
osteointegration. 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 imcromovement during the healing
phase. Such surfaces were also reported to allow the
successful placement of shorter length implants when
compared with noncoated implants. The basic theory
was based on increased area for bone contact.
134. 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. In addition,
the present technology allows metallurgic bonding of
coatings and a high resistance against mechanical
separation of the coating with many coating test values
exceeding the published standard requirements.
135. 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 than 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.
136. 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 m,
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.
138. 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.108,191 One of the concerns
regarding CPC coatings is the strength of the bond
between the CPC and the metallic substrate. Investigate
ion-beam sputtering coating techniques for CPC or CPC-
like nonresorbable coatings to varied substrates appear
to produce dense, more tenacious and thinner
139. 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.
140. 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. Animal studies also show reductions in coating
thickness after in vivo function. 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.
141. The concerns related to calcium phosphate coatings
have focused on (1) the biomechanical stability of the
coatings and coating-to substrate interface under in
vivo conditions of cyclic loading, and (2) the biochemical
stability of these coatings and interfaces within the
gingival sulcus (especially in the presence of
inflammation or infection) and during enzymatic process
associated with osteoclasis remodeling of the bone-to-
coating interfacial zones.
142. Other Surface Modifications
Surface modification methods include controlled
chemical reactions with nitrogen or other elements or
surface ion implantation procedures. The reaction of
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.
143. 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.
144. 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).
145. 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 range
146. 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
147. 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.
As intrinsically 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.
148. 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.
149. 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.
150. 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. Preliminary studies by Schmidt and
Grabovski et al. showed modified fibroblast adhesion on
nitrogen and caron-ion implanted titanium. A general
rule has been that cleaner is better.
151. 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.
152. Metal or alloy constituents, inorganic and organic
particles, corrosion products, polymers, and precipitates
can be absorbed at the surface throughout the
manufacturing, polishing, cleaning, sterilization,
packaging, and storaging processes. Baier and Meyer
correlated the usual type of contaminant found in
relation to the sterilization technique used. 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.
153. 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.
154. 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 and their action
a films 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.
155. 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
156. 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.
157. 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.
158. 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.
159. REVIEW OF LITERATURE
Gluszek et al (1990) conducted a study wherein Steel 316L was
coated with titanium or titanium nitride by ion plating. The
tightness of the coatings was examined electro-chemically. The
galvanic effects for the galvanic couples steel-titanium, steel-
titanium-coated steel and steel-titanium nitride-coated steel were
studied. It was found that both titanium and titanium nitride
coatings were non-porous in Ringer's solution; titanium served as
an anode in the couple steel-titanium; it was oxidized according to
the logarithmic law. For the other two couples, the coatings were
the cathodes. The rate of dissolution of steel in these couples,
was however, smaller than expected, owing to a strong
polarization of the coatings. The potential of the couple was
similar to that of steel.
160. Denissen et al (1996) Calcium phosphate ceramic coatings with a
hydroxyapatite chemistry applied on the surface of dental implants
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. This study compared
the behavior in bone of newly developed fluorapatite and heat-treated
hydroxyapatite coatings, with the clinically used hydroxyapatite coatings
used as controls in experimental studies in dogs. The biologic responses
to fluorapatite and heat-treated hydroxyapatite coatings were the same
as those to hydroxyapatite coatings, and bone condensation around all
coatings was histologically evident. However, the coating thickness of
the fluorapatite and heat-treated hydroxyapatite coatings remained
stable with only minor changes during the observation period of 24
months.
161. Cross-Poline GN et al (1997) compared the surface roughness
produced by various implant curets on titanium implant abutment
surfaces. Each of six titanium implants was divided into four quadrants,
three experimental and an untreated control surface. 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.
162. Augthun M et al (1998) examined The effect of specific cleaning
procedures on the surfaces of 3 implant types with different coatings and
shapes (plasma sprayed [PS]; hydroxyapatite coated [HA] implants; and
smooth titanium surface screws) using a scanning electron microscope.
Each implant was treated for 60 seconds per instrument with one of 6
different hygiene measures: plastic curet, metal curet, diamond polishing
device, ultrasonic scaler, air-powder-water spray with sodium
hydrocarbonate solution, and chlorhexidine 0.1% solution rinse. The air-
powder-abrasive system, chlorhexidine rinse, and curettage with a plastic
instrument caused little or no surface damage in all but the hydroxyapatite-
coated fixtures. Therefore, these 3 methods were tested to determine their
cleaning efficacy in a second clinical study, which did not include the HA-
coated fixture. 2 fixtures on each side were examined in each patient. The
examination revealed that only the sodium hydrocarbonate spray yielded a
clean fixture without damage to the implant surface. In a third stage, which
imitated the clinical procedure of the second approach, the cell growth of
mouse-fibroblasts on implant surfaces was examined after cleaning the
surface with plastic scaler and the air-abrasive system, which represents
the least damaging and most effective methods. In contrast to the implant
surfaces treated with plastic scalers, mostly vital cells were found on
implants sprayed with the air-abrasive system.
163. 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 for the electrochemical tests,
in a pH range of 6.15 to 3.0. The mixed potential theory was
applied to predict couple potentials and couple currents. Thus (a)
with and without fluoride ions, galvanic currents are weak within a
pH range of 6.15 to 3.5; (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 of 6.15 to 3.5; (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.
164. Sawase et al (2001) The surface oxide layer of titanium plays a
decisive role in determining biocompatibility. However, there are some
reports demonstrating that the natural oxide film may not be sufficiently
protective in the aggressive biologic environment. The goal of this 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. Improved corrosion
resistance was demonstrated by the potentiodynamic polarization
measurements. Light microscopic histomorphometry showed that all
implants were in contact with bone and had some proportion of bone
within the threads at 4 weeks; however, there were no significant
differences compared with as-machined controls. 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.
