1. INTRODUCTION

1.1. MATERIALS CHEMISTRY

Chemistry is the study of the composition, structure and properties of

substances and the transformation by ,wich substances are changed to other

substances [I). Chemistry has direct relationships with all the material aspect of the

world Everything in this world comes under the title chemistry, as every part of

matter is made up of atoms/molecules. Growing acceptance in recent years of the

importance of chemistry to material preparation and processing bas resulted in the

recognition of materials chemistry as a distinct sub discipline of chemistry. The

material world bas influence on the cultural,, socio-economic,, demographic and

geographic development of the society [2].

A material is something that has properties ,,ib.ich give a particular useful

application, either structural, as \'ilrith building material or functional, as ,vith

materials used to make de\-"ices ( electronic, optical or magnetic). Chemistry is

important both in the development of ne,v and novel materials and in the improved

processing of established ones. Materials chemistry in general can be defined as

the chemical science that deals with the preparation, processing and analysis of

solid state materials [3].

Human civilization is defined by the man's relationship with materials­

namely Stone Age, Bronze Age,. Iron Age etc. Utilization of materials

distinguished humans from animals. Effective integration of material science \Vith

its macroscopic· perspective and chemistry which focuses on atomic and molecular

level interactions could pro,,.1de opportunity to understand and control the

2

fundamental connections between structure and functions from molecular level to a . .o

macroscopic level. This could lead to improved composition, structure or synthetic

methods and enable the development of new types of advanced performance

materials having superior properties and performance.

Chemistry is involved in the generation and processing of materials. Basic

chemistry involved in the generation of metals is historical and largely empirical.

Metals are derived from their ores by chemical reduction of their oxides or silicates

or by oxidation of their sulphides. Metals are developed from rocks, minerals

directly and later developed through experimentation. An important example of

this is the development of ceramics, such as pottery, glass etc. and later hydraulic

cement. The term ceramics include all inorganic engineering materials other than

metals and semiconductors. The ceramic processing involves extensive purification

and processing before their consumption as powders. The ceramic processing

methods include spray drying, vapour phase synthesis, sol-gel processing,

chemical vapour deposition and infiltration and polymer precursor pyrolysis. The

oxidation, doping, patterning and etching involved in ceramic processing is largely

and entirely chemical in nature [4].

The development of human civilization has a direct relationship with

material processing and consumption. Therefore the development of materials

chemistry is very much significant.

1.2. DEVELOPMENTS IN MATERIALS CHEMISTRY

The earliest ages of human civilization are classified by the key materials in

use at that time: stone, bronze and iron. The later industrial age was characterized

3

by the production of materials on a large scale for commercial purposes. Metals

have been among the most influential materials in the early development of

civilization and chemistry has been intimately involved in their production. All

early materials were developed by an empirical process, as the micro structural

consequences of thermal and physical processing became recognized, various

refinements in materials processing has developed. Advances in chemical

understanding, together with advancement in analytical methods led to

improvements in materials technologies.

Natural organic polymers, in the form of wood and other plant fibers, were

among the first materials to be used by humans. The conversion of wood into paper

was a contribution of the Chineese that was transmitted westward by the Arabs in

the period AD 750-800 [5]. Now the role of paper is gradually transmitting into

plastics, ceramics and semiconductors in the form of CDs, magnetic recording

media and integrated circuits. The first synthetic natural polymeric material used

by humans is silk and cobwebs. The first impact of chemistry in the development

of synthetic polymers came in 1839, when Charles Goodyear heated natural rubber

with sulphur. Another polymeric material cellulose nitrate, developed became the

basis for guncotton and cellophane film. The first fully synthetic polymer was

produced in the early twentieth century by Leo Baekeland with the production of

Bakelite, a polymer from phenol and formaldehyde. The revolution in synthetic

polymeric materials led to the production of nylon, rayon, teflon and lexan.

The materials were classified based on their typical chemical constitution

and their typical physical properties. Alloys are combinations of one or more

metallic materials. Ceramics are compounds formed between metallic and non

4

metallic elements e. g., oxides, sulphides, nitrides and carbides and also natural

minerals of the earth as well as glass ceramics. Polymers are comprised of

macromolecules that range from linear polymers having amorphous structure with

extensively cross linked networks. Composites are combinations of different

materials e. g., fiber glass in which glass fibers are embedded within a polymeric

matrix. Composite materials have the advantages of light weight and strength. The

total material cycle is exhibited in Fig. 1.1.

E:i.."tract Refine

ProceEOs

Bulle ]yfa terla Is

Chemica s Metals

Paper Cewent Fibers

Process

Arena of mineral lllld asui crdtural

��-� sciences and

Arena of materials sciences

Recycle and eugiueeriug

Engineerin!!

Dispose

Performance �-� Service

Use

Fig. 1.1. The total material cycle

Engineering materials

From the total materials cycle it can be recognized that the entire process of

materials fabrication is intimately interconnected with both the materials origin and

its ultimate disposal or recycling. The total materials cycle is used to indicate the

intimate relationship between the continuing need for new materials and the effects

of materials development on the environment and on the diminishing supply of

basic raw materials [6]. The processes included in the cycle are basically chemical

5

processes and the role of chemists in the solution of fu�damental materials science

problems is becoming important with time.

New materials and processes will continue to be needed and chemistry will

continue to play a large role in their development. The nanometer size scale

materials are developing with a large pace. Nanomaterial is one which has

nanometer (typically 1 to 100 nm) size range. The materials in the nanosize regime

exhibit size-dependent properties. Nanosized particles find wide use in molecular

storage devices as quantum dots, which exhibit a size quantization effect in at least

one dimension. Molecular storage devices might employ molecular/atomic wires.

Hence connections might be drawn from natural systems and this brings the field

of biomaterials.

Biomaterials include everything from materials produced by biological

organisms or as a result of biological processes, to synthetic materials whose

design form or construction is inspired by analogies to biological systems. One

important aspect of biomaterials is the need for temporary or permanent

replacement of parts of the body. The search for such materials is a major area of

current research. The use of biomaterials and its wide range of application and

properties are described in the following sections.

1.3. BIOMATERIALS

Biomaterial is a substance that is used in prosthesis or in medical devices

designed for contact with the living body for an intended application and for

intended time period. Biomaterial can be defined as "a non-variable material used

in a medical device intended to interact with biological systems" [7]. Materials in

6

surgical implants & medical devices are non toxic, not cause any deleterious

effects. It should have the ability to perform with an appropriate host response in a

specific application [7, 8]. Biomaterials are pharmacologically inert substances. A

biomaterial must be in contact with living tissues or body fluids resulting in an

interface between living and non living substances.

1.3.1. Historic developments in biomaterials

Table 1.1. lists out the historic developments in the field of biomaterials [9,

10]. The earliest operations performed were by Hindu surgeons for restoration of

missing parts. Sushrutha, in about 600 BC, repaired an injured nose. This

technique for nose reconstruction migrated from east to west. Around 1430, the

Brancas, a family of civilian laymen, perfected the Italian method for nose

reconstruction by using skin flap taken off from the arm. In the 19th century, Von

Graefe and Dieffenbach recorded several techniques for reconstruction of missing

parts. The major trajedies during the First World War pioneered newer methods

Table 1.1. Major historical developments of biomaterials

Year Author Activity

600BC

Late 18th -19th

century

1860-1870

1893-1912

Sushruta Samhitha

J. Lister

W.A. Lane

Nose reconstruction

Various metal devices to fix

fractures, wires and pins made of

Fe, Au, Ag and Pt.

Asceptic surgical techniques

developed

Steel screws and plates for fracture

7

fixation

1912 W. D. Sherman Vanadium steel plate, first alloy

developed exclusively for medical

use, less stress concentration and

corrosion

1926 E. W. Hey- Groves Used Carpenters screw for femoral

neck fracture fixation

1926 M.Z. Large 18-8 s Mo (2-4% Mo) stainless steel

for greater corrosion resistance than

18-8 stainless steel.

1931 M. N. Smith Petersen Designed first femoral neck fracture

fixation nail made originally from

SS, later changed to vitallium®

1936 C. S. Venable, W. G. Vitallium® ( developed in 1929; 19

Stuck w/o Cr-9 w/o Ni stainless steel)

1938 P. Wiles First total hip replacement

1940's M. J. Dorzeee, A. Acrylics for corneal replacement

Franceschetti

1944 W.J. Kolff Hemodialyser

1946 J. Judet and R. Judet First bio mechanically designed hip

prosthesis. First plastics used in joint

replacement.

1952 A. B. Voorhees, First blood vessel replacement made

1953

1958

1958

1960

1980's

A. J aretzta,

A. H. Blackmore

A. Kantrowitz

S. Furman,

G. Robinson,

J. Charnley

S. Furman, G.

Robinson

A. Starr,

M. I. Edwards

W. J. Kolff et al.

of cloth

Intra aortic balloon pumping

First use of acrylic bone cement in

total hip replacements

First successful direct stimulation of

heart

Heart valve

Artificial heart

8

of wound closure and tissue transfer. The earliest written record of metals m

surgical procedures is from the year 1565. The modem implant developments,

which centered on repairing long bones and joints, began at the end of the 19th

century. With the beginning of the plastic industry in the 1930's, the polymers find

a variety of applications. The second world war gave pace to the.implant industry.

Heart valve implantation was possible only after the development of open heart

surgery.

1.3.2. Classification

Medical biomaterials are classified into five classes - metals, ceramics,

composites, synthetic polymers and biopolymers. Tablel.2. depict the variety of

materials used in implants.

Materials

Polymers

Polyolefins

Polyesters

Polyamides

Polyurethane

Polyacetals

Polyether

Silicone-rubber

Metals

Stainless steel

Titanium

alloys

Cobalt-

chromium-

alloys

9

Table 1.2. The materials used in implants

Advantages

Low density

Easy to

fabricate

High impact

strength

High

resistance to

wear, ductile,

absorption of

high strain

energy

Disadvantages

Low mechanical

strength, additives,

Common applications

Cardiovascular,

maxillofacial, soft

oligomers may cause skeletal tissue such as

tissue reactions

Low

biocompatibility,

corrosion in

physiological

environment,

mismatch for

mechanical

properties with soft

tendon, ligament, space

filling devices, dental

implants, bone cement,

lens and middle ear

prosthesis,

adhesives,

delivery systems

Orthopaedic load

tissue

drug

bearing and fixation

devices, dental

implants

10

convective tissues

Pt, Pt-Ir alloys High Low mechanical Neuro muscular

conductivity strength High cost stimulation

Ceramics

Alumina Good Undesirable surface Hip and knee

Zirconia biocompatibilit properties, special prosthesis, dental

y, inert, techniques are implants improving

corrosion needed for material biocompatibility

resistance, fabrication

high tensile

strength

Calcium Degradation not Temporary support

phosphates Biodegradable controllable assist regeneration of

natural tissues

The different classes of biomaterials are the following

a) Metals

Bioactive metallic materials are largely used for implantation purpose.

They are employed for two primary purposes. They are used as prosthesis to

replace a portion of the body such as joints, living bones and skull plates. They are

also used as fixation devices to stabilize broken bones and other tissues while

normal healing proceeds. Some of the engineering materials presently used for

implants include stainless steel (SS), Co and Ti based alloys and conducting metals

11

such as Pt and Ir. Most metals used for manufacturing implants such as Fe, Cr, Co,

Ni, Ti, Ta, Mo & W can be tolerated by the body only in minute amounts. The high

modulus and yield strength together with the ductility of metals make them

suitable for bearing large loads without leading to large deformations and

permanent dimensional change.

b) Ceramics

The disability of metals to bond with bone resulted in the popularity of

ceramics [11-13]. The ceramics used in implantation and clinical purpose include

alumina, partially stabilized zirconia (PSZ, Y-TZP and Mg-PSZ), bioglass-glass

ceramics, calcium phosphates and crystalline and glassy forms of carbon and its

compounds [14]. Alumina and zirconia ceramics have high strength and

biocompatibility suitable for load bearing applications (15, 16]. Scandia 1s a

potential biomaterial. It had no cytotoxic and deleterious effect on cell metabolism

[17]. The ceramic materials used in reconstructive bone surgery are

osteoconductive (3]. Alumina ceramics were introduced nearly 30 years ago as a

candidate material for bearing surfaces in hip replacements. Mg-PSZ zirconia

ceramics were introduced in the 80's. It was replaced by Y-TZP zirconia ceramics

which offers better mechanical strength. A novel biocompatible ceramic is now

developed which is a combination of hard alumina and tough zirconia. This offers

the option to improve the fracture toughness and the mechanical strength and to

design ceramic components for total hip replacements [ 18].

c) Composites

Composhe materials are a mixture of two or more phases bonded together

so that stress transfer occurs across the phase boundary. Typically composite

12

materials are designed to provide a combination of properties that cannot be

achieved with single phase material. With the use of composites there 1s a

possibility to produce light weight, high strength structural member with

anisotropic properties similar to those of natural bone. Composite materials exhibit

high bonding strength [19, 20]. Poly (E-caprolacton)/silicahybrid composite is used

as bone substitute [21, 22]. TiOi/ultra high molecular weight polyethylene is used

for bone repairing applications [23, 24].

d) Synthetic polymers

Polymers are very large molecules made by the repetition of small, simple

chemical units termed monomers. The repetition of monomers appears as straight

or branched chains. The structure of polymer determines their chemical and

physical properties. Biomedical polymers can be classified into elastomers or

plastics. Elastomers are able to withstand large deformations and return to their

original dimensions after releasing the stretching force. Elastomers include butyl

rubber, chloro sulfonated polyethylene (Hypalon®), epichlorohydrin rubber

(Hydrin®), poly urethane, natural rubber and silicone rubber. But plastics are rigid

polymers. Plastics can be classified into thermoplastics and thermosetting plastics.

Thermoplastic polymers can be melted, reshaped and reformed. Thermosetting

plastics cannot be remelted and reused. Thermoplastic polymer used as

biomaterials include polyolefins, teflon® (fluorinated hydrocarbons),

polymethylmethacrylate (PMMA), polyvinylchloride (PVC), polycarbonate, nylon,

polyester etc. Hydrogels are the first biomaterials used in the human body for soft

implants, i. e., contact lenses [25]. They are water soluble polymeric material and

13

maintain a distinct 3D structure. Ethylene-vinyl alcohol co-polymer could be

modified with a silane coupling agent and is suitable for implantation purpose [26].

e) Biopolymers

Bioploymers are polymers formed in nature during the growth cycle of all

organisms. They are formed within the cells by complex metabolic processes. A

variety of biopolymers find application as biomaterials. The prominent among

them are collagens, mucopolysaccharides, chitin, cellulose and its derivatives.

Collagens are major animal structural proteins and are widely used in a variety of

forms such as solution, gel, fibers, membranes, sponge and tubing for a large

number of biomedical applications including drug delivery systems, sutures,

vessels, valves, corneal prosthesis, wound dressing, cartilage substitute and in

dental applications. Cellulose and its derivatives are mainly employed in the

fabrication of membranes utilized in hemodialysis machines. Heparin and

sulphated mucopolysaccharides find application for improving blood compatibility

of other materials. Chitosan and chitin materials are emerging as biomaterials with

wide applications. A large number of carbohydrate molecules find application as

biopolymers.

1.4. BIOMATERIALS IN ORTHOPAEDICS

Biomaterials are largely needed to replace, repair or regenerate injured or

diseased bone. Bone is among the most frequently transplanted tissues [3]. Human

bone is a dynamic living tissue and changes throughout life [27]. Like other

connective tissues of the body bone comprises cells embedded in an abundant extra

cellular matrix mineralized to bestow unique physiological functions. Human bone

14

is made up of composite material- collagen and calcium phosphate mineral. Adult

bone comprises 10% water, 30-40% collagen, approximately 60-70% mineral

deposit [17]. The bone mineral calcium phosphate is hydroxyapatite (HA)

containing carbonate and small amounts of sodium, magnesium, fluoride and other

trace elements [28]. HA have the chemical composition CaIQ(P04)6(OHh with a

Ca/P ratio of 1.67. Bone is unique among the tissues of the body in the level of its

resistance to compressive forces. The skeleton has numerous functions, but the

main function is to provide shape and size to the body and to give mechanical

strength. Hence orthopaedic implants are designed in a manner that it can bear

load. Also the interfacial bonding between the mineral and organic constituents is

concerned while designing orthopaedic implants. Fig.1.2. depicts the natural and

artificial hip implant. The metallic implant is fixed in the body with the help of a

bone joining ceramic material and the brown coloured portion is the cement

material.

Fig. 1.2. Natural and artificial hip joints

15

F. D. A. Fed. Register defines an implant as "a device that is placed into a

surgically or naturally formed cavity of the human body if it is intended to remain

there for a period of 30 days or more". Approximately 200,000 hips are replaced in

US per year. As the life expectancy and activity levels of the ageing population

increases this number will continue to grow [29, 30]. Hip fracture is most frequent

in elder persons. For those over 80 years hip fracture is very common. Globally the

number of fracture may rise to 2.6 million by the year 2025 and between 7.3 - 21.3

million by 2050 [31]. The annual value of implant market is $ 1 billion. Numerous

implants are used for the treatment of hip fracture. More important to the choice of

implant is the surgical technique and efficiency [32]. After 15 years of post

operative use 15-25% of devices failed and requiring costly and traumatic revision

surgery. Implant loosening in 20% case by 5 years post operative cases and 30%

by ten years [33]. This is mainly due to the inadequate bonding between implant __

and bone. An ideal bone substrate must be able to function as a scaffold and to

participate in the formation of new bone. The implant must be biocompatible and

bioreabsorbable so that it may be completely substituted by newly formed bone. To

meet these requirements biologically compatible materials, calcium phosphates­

based bioactives and glass-based bioactives-have been developed [3]. Various

engineering materials are used in orthopaedic surgery as replacement materials of

bone. The materials include metals and ceramics.

1.4.1. Metals

Bioactive metallic materials which can directly bond to living bone are

largely needed in. dental and orthopaedic implants for better long-term results. One

of the methods to increase the bioactivity of metal is providing an apatite layer on

16

the surface of the metal. Metals were used in surgical procedures from the year

1565. With the invention of X-ray by Roentgen in 1875, the fractured bones were

visualized and then metals were largely used in surgical procedures. The high

modulus and yield point coupled with ductility of metals make them suitable for

bearing large loads without leading to large deformations and permanent

dimensional changes. Metallic implants are used as prosthesis to replace a portion

of the body such as joints, long bones and skull plates and also as a fixation device

used to stabilize broken bones and other tissues while normal healing proceeds.

The commonly used metallic implants are depicted in Fig. 1.3. Metals are used as

screws, nails, bone plates and hip joints.

Fig. 1.3. Commonly used metallic implants

17

Most metals used for manufacturing implants can be tolerated by the body

in minute amounts, but cannot be in large amounts. Some of the engineering

materials presently used for implants include Titanium (Ti), stainless steels (SS),

Co-based alloys, Ta, Pt and Ir metals. Titanium has very high biocompatibility and

is widely used for biomedical applications [34, 35].

1.4.1.1. Titanium

Titanium (Ti) and its alloys find wide range of applications in aerospace,

marine and clinical industry to medicine due to its remarkable mechanical and

electrochemical properties. The oxide film is responsible for the corrosion

resistance of titanium [36]. Titanium is highly bioactive and bums violently in dry

chlorine, but it has extremely protective passive film in natural environments

including sijlt solutions, most oxidizing acids and many organic compounds [37-

39]. Titanium and its alloys are most biocompatible metallic biomaterials [13-16,

40, 41]. Titanium have elastic stimulus similar to bone [42, 43]. Titanium is

biologically inert and highly corrosion resistant. This nature is due to the

spontaneous formation of Ti02 thin film on its surface in air and in electrolytes

[44-48]. These Ti02 thin films have good biocompatibility [49], but not bioactive

enough to induce calcium phosphate formation [50].

Surface native titanium oxide consists of non stoichiometric titanium oxide

and display amorphous phase of low crystallinity. After implantation for a long

period, Ti implants cannot bond directly to living bone. Ti ions releases into the

bio-liquid, in the metal-tissue interaction studies, a high concentration of Ti have

been detected both in osseous tissue and surrounding organs after three years.

From immersion tests it was found that anatase Ti02 exhibit excellent bio activity

18

[51]. Ti-OH groups of anatase structures are effective for the apatite nucleation

[52] whereas those of amorphous structure and Na2Ti5011 crystal are not effective.

In vitro apatite deposition is induced on anatase structure by the Ti-OH groups

[53]. The commercially pure Ti disks after polishing were heated at 550, 600, 650

and 700°C by thermal oxidation technique. Amorphous Ti02 layer have

heterogenous structure. Ti annealed at 600°C and below in air and at 600°C in

argon has high calcium phosphate (CaP) forming ability. The rutile Ti02 has no

CaP precipitation [54].

The new surface formed on Ti improves the performance of the implant by

accepting the device by the body, long term integrity and stability of the

biomaterial/body interface [55]. Anodically oxidized Ti plates were bonded

directly with no intervening soft tissue layer and no breakage of the anodic

oxidized Ti layer is found upon histological and morphological investigation [56].

Ti02 layer formed on the Ti surface is very beneficial for improving further

adhesion of the coating [57]. Titania with specific structures of anatase and rutile

was found to induce apatite formation in vitro. Anodic oxidation in H2S04 solution

could form anatase and rutile titania on metal surface. The induction time of apatite

formation decreased with increased amount of either anatase/rutile by anodic

oxidation [58]. A titanium oxide gel electrochemically prepared on Ti surfaces

induces bio activity [59, 60]. Cathodic reactions in electrolytic cell produce

negative charges on metal surface. This charge formation is suitable for activating

Ti [61].

Alkaline treatment proposed by T." Kokubo et. al is the most effective

method for increasing the bio activity of Ti. In alkaline treatment Ti was immersed

19

in NaOH at 60°C for 24 h followed by thermal treatment at 600°C for 1 h to

produce a bioactive sodium titanate layer [62, 63]. This bioactive layer is

responsible for apatite nucleation [63-68]. In alkaline treatment the surface get a lot

of negative charges. The titanate/titania gel layers rich in Ti-OH or Ti-0 groups

served as the sites for nucleation induction, both in vivo and in vitro and then the

calcium phosphate crystals grow spontaneously on Ti surface.

Osteoblast adhesion increased on HA coatings after hydrothermal

treatment. Ti was coated with HA, Ti02 and poly(dl-lactic-glycolic acid, PLGA)

using sol-gel processing. The calcium phosphate coatings deposited were subjected

to hydrothermal treatment [69]. The biocompatibility of Ti is very high. Hence Ti

metal is commonly used as implant material.

1.4.1.2. Ti Alloys Ti & its alloys are widely recognized as useful structural materials. Certain

Ti alloys showing exceptional strength to weight ratio and good mechanical

properties. Ti6Al4V is the widely used alloy to manufacture implants. The main

alloying elements are aluminium (5.5-6.5%) and vanadium (3.5-4.5%). Ti is a light

metal (density 4.505 g/cm3 at 25°C). Since Al is a lighter element and V barely

heavier than Ti, the density of Ti6Al4V alloy is very similar to pure Ti. Ti alloys

have lower Young's modulus and it gives great flexibility to orthopaedic devices.

Ti alloys have high specific strength and it outperform any implant material in this

matter. But the high coefficient of friction causes wear and releases wear particles.

20

The extensive biomedical application of Ti6Al4V is due to good

mechanical property and ability for osteointegration [ 40]. Very stable and adherent

oxidati ')fl film is formed on the alloy surface and it result in an excellent corrosion

resistance and passivity [70-74]. Hydroxyapatite coating is developed on Ti6Al4V

alloy and it increases the corrosion resistance [75, 76].

1.4.1.3. Stainless Steel (SS)

The austenitic stainless steel, 316 & 316L are widely used for implant

fabrication. SS has a low content of impurities and a passivation is suitable for

implantation in the human body. SS is a predominant implant alloy due to its ease

of fabrication, desirable mechanical properties and corrosion behaviour. SS is first

introduced in surgery in 1926. In 1943, 302 SS is recommended for bone fixation.

In the 1950' s 316L SS was developed by the reduction of maximum carbon

content from 0.08% to 0.03% for better corrosion resistance. The prime alloying

elements are Fe, Cr, Ni, Mo & Mn, Cr on the surface of steel provide excellent

corrosion resistance. Mo in steel enhances the resistance to pitting corrosion.

Table 1.3. Composition of Stainless Steel (Balance o/o iron)

Type ofSS

301

304

316, 18-8s Mo

316L

430F

%C %Cr

0.15 16-18

0.07 17-19

0.07 16-18

0.03 16-18

0.08 16-18

%Ni %Mn % other elements

6-8 2.0 1.0 Si

8-11 2.0 1-Si

10-14 2.0 2-3 Mo, 1.0 Si

10-14 2.0 2-3 Mo, 0.75 Si

1.0-1.5 1.5 1.0 Si, 0-6 Mo

21

3161 SS may corrode inside the body under certain circumstances in a highly stressed and oxygen depleted region. Hence SS is suitable to use only in temporary implant devices, such as fracture plates, screws and hip nails. Due to the low biocompatibility and fatigue life SS is inferior to other super alloys. SS lead to interface corrosion. Surface finish can give significant influence on the corrosion resistance of SS. The surface oxide layers increases the corrosion resistance of SS [80]. Literature reports that HA coating is developed on 3161 SS [78].

Passive film is formed on SS with the addition of complexing agent. The film contains two predominant oxides-chromium and iron oxides. Also oxides of alloying elements of Ni & Mo induce passivity. The strong enrichment of the passive films is responsible for the corrosion resistance of orthopaedic SS in physiological solution. An artificially induced passive layer on 3161 SS surface prior to HA coating improved the nature of the implant' s corrosion resistance. HN03 treatments on 3161 SS provide a passive oxide layer [79]. Alumina, magnesia-stabilized zirconia (MSZ) and yttria stabilized zirconia (YSZ)-of different thickness over 3161 SS are produced by air plasma spray technique. MSZ coating (thickness 150 µm) increases the breakdown potential value when immersed in HBSS of pH 7.4 at 30°C and it improves the corrosion resistance of 316 1 SS considerably [80]. HA-3161 SS fiber composites were prepared. Micro cracking took place invariably in the HA matrices but near and around the 3161 fibers. Such patterns of micro cracks resulted from the thermal residual stresses developed during cooling from the high temperatures and the intrinsic low mechanical strength of HA ceramics [81]. Calcium phosphate layer can get

22

deposited on 316L SS when immersed in Simulated Body Fluid (S. B. F) after

chemical pre-treatment and thermal treatment [82].

1.4.1.4. Co-Cr-Mo alloy

There are basically two types of cobalt chromium alloys. One is the Co-Cr­

Mo alloy, which is usually used to cast a product and the other is the Co-Ni-Cr-Mo

alloy, which is usually wrought by forging. The castable Co-Cr-Mo alloy has been

used for many decades in dentistry and in making artificial joints. The Co-Ni-Cr­

Mo alloy is used for making stems for prosthesis for heavily loaded joints such as

knee & hip. Cobalt based alloys are highly resistant to corrosion and especially to

attack by chloride within crevice. These alloys are also quite resistant to fatigue

and to cracking caused by corrosion, and they are not brittle, since they have a

minimum of 8% elongation. The superior fatigue and ultimate tensile strength of

the wrought Co-Ni-Cr-Mo alloy make it suitable for the applications which require

long service life without fracture or fatigue. It is used for the stems of the hip joint

prosthesis. The modulus of elasticity of the Co-Cr-Mo alloys does not change with

the changes in their ultimate strength. The values are higher than other materials

such as SS. This may have some implications of load transfer modes to the bone in

artificial joint replacements, although the effect of the increased modulus on the

fixation and longevity of the implants is not clear.

The two basic elements of Co based alloys form a solid solution of up to 65

wt% Co and 35 wt% Cr results in higher strength. During 1930' s Co-Cr-W alloys

are widely used for the preparation of metallic dental castings as alternatives to

gold alloy. Many· of the alloys used in dentistry and surgery, based on the Co-Cr

system contain additional elements such as carbon, molybdenum, nickel, tungsten

23 and iron. The elastic modulus vanes from 185-250 GN/m2 depending on the composition, being equivalent to that of 316L SS and twice as that of Ti. Co-Cr-Mo alloys used in artificial joint and bone had high wear resistance [83]. 1.4.2. Ceramics Ceramics are classified into three types 1) inert 2) surface reactive and 3) completely resorbable. Inert ceramics include carbons, alumina, zirconia etc. Glass ceramics are surface reactive ceramics. Calcium phosphate ceramics are completely resorbable ceramics. 1.4.2.1. Carbons

The carbons are inert ceramic materials, which exhibit varied and unique properties that are not found in any other materials. In the quasi-crystalline forms, the degree. of perfection of the crystalline structure and the morphological arrangements of the crystallites and pores are important in determining the properties of carbons.· Currently all the carbons used in medical devices have the quasi-crystalline turbo static structure.

The covalent bond that binds the atom within the hexagonal layers is responsible for the high strength of the material. The weak van der Waals bonding causes low stiffness to carbon. The weak bonding between the layers permits large shear strains at low stresses. The elastic modulus of near 20 G Pa and density range 1.5 g/cm3 to 2.29 g/cm3 of carbons is in close proximity with those of bone.

Carbon coatings find wide applications in heart valves, blood vessel grafts, percutaneous devices because of its exceptional compatibility with soft tissues and blood. This causes reduction in critical surface tension and blood adhesion. Platelet adhesion and activation is found to be least with carbon coated surfaces. Hence

24

ultra low temperature isotropic carbon (UL TI, i. e., carbon vapour deposited at low

temperature) coated valves is most widely used. Carbon does not provoke an

inflammatory response in adjacent tissues and no foreign body reactions to the

material have been observed. Bone and soft tissues are much more tolerant to

carbon than other materials. A thin sheath like capsule is formed around the carbon

coated implant, which isolate it from surrounding tissue.

1.4.2.2. Alumina

Alumina is an inert ceramic. It has high corrosion and wear resistance.

Alumina is obtained by the calcinations of aluminumtrihydrate. High density

alumina is used in load bearing hip prosthesis and dental implants because of its

combination of excellent corrosion resistance, good biocompatibility, high wear

resistance .and reasonable strength. Most alumina devices are fine grained

polycrystalline oc-alumina. Strength, fatigue resistance and fracture toughness of

polycrystalline oc-alumina are function of grain size and purity. An increase in

grain size from 4 µm to 7 µm can decrease the mechanical strength by 20%.

Alumina is used for the replacement of hip and knee joints, shoulders, radius,

vertebra and ankle joint prosthesis. Alumina is not cytotoxic and shows no

inflammatory or progressive fibrotic reactions. HA coating is provided on alumina

to enhance its bone bonding ability [84]. Alumina reinforced HA porous implants

have higher strength than HA porous implants and exhibited similar bio activity

and osteoconduction property to the HA porous implants [85]. Reducing the grain

size of alumina improved the toughness of the ceramic [86]. But alumina is inferior

in calcification than Ti02 surface when immersed in S. B. F [87].

25

1.4.2.3. Zirconia

At room temperature zirconia has monoclinic structure. Upon heating it transforms to a tetragonal phase at 100-1100°C and cubic phase at around 2000°C. Yttria oxide (Y203) stabilizes the tetragonal phase; upon cooling the tetragonal crystals made of Zr02-Y 203 can be maintained in meta stable state and not transform in a monoclinic structure. Young's modulus is half of that of alumina, while the bonding strength and fracture toughness is greater. The increased mechanical properties may allow for smaller diameter femoral heads to be used compared to alumina. The wear resistance is a function of fine grain size.

Yttria stabilized zirconia have excellent biocompatibility and wear properties and are used for orthopaedic applications [14]. HA-Zr02 composite coating is produced by plasma spray method [88, 89]. Zirconia toughned apatite implant is used for orthopaedic applications [90]. 1.4.2.4. Glass ceramics Glass ceramics is a surface reactive ceramic and it can achieve a controlled surface reactivity that will induce a direct chemical bond between the implant and the surrounding tissue. It is used as bone substitute [91]. Glass ceramics are polycrystalline ceramic made by controlled crystallization of glass. They are manufactured by controlled nucleation & growth of crystals of small ( <l µm) uniform size. About 1012-1015 nuclei per cubic centimeter are required to achievesmall crystals. In addition to the metals of platinum group, metal oxide such as Ti02, Zr02 & P2Qs are widely used for nucleation.

26

Bioactive glass is developed in l 970's. Bioglass implants have several

advantages. The mechanical strength is not limited because the material used need

not be porous. This can be applied as a coating to high strength SS, Co-Cr alloys or

Ah03 providing a combination of high mechanical stability along with surface

biocompatible properties. The surface reactive compounds respond to the local pH

changes by releasing ca2+, Na+ and K+ ions. The bioglass ceramic containing less

reactive fluorides acquire a fibrous capsule when implanted in rat femurs.

The essential condition for glasses and glass ceramics to bond to living

bone is the formation of an apatite layer on their surfaces in the body. Apatite

forming ability increases on increasing CaO content, although the tensile strength

and Young's modulus decreased [92]. Hydrated silica formed on the surface of

these materials in the body plays an important role in forming the surface apatite

layer [93]. Tissue bonding occurs on the surface of bioactive glasses [94]. Also

biomimetic apatite deposition takes place on the surface of bioactive CaP invert

glasses [95]. A 3-D ordered macro porous sol-gel bioactive glasses have bio

activity and degradability [96]. Bioactive glasses are now the emerging field in the

case of orthopaedic implants.

1.4.2.5. Calcium phosphates

Calcium phosphate is a resorbable ceramic. Hydroxyapatite (HA) is the

important calcium phosphate group material having the composition [Ca10(P04)6

(OH)i]. It is the natural bone material and is the major inorganic component of

vertebrate bone and teeth. HA crystallizes into the hexagonal rhombic prism. The

unit cell has dimensions of a = 0.9432 nm and c = 0. 6881 nm. The ideal Ca/P ratio

27

is 10/6. Its density is 3.219 g/mL. HA has a higher elastic modulus than other

mineralized tissues.

HA can be made more bioactive by substitution. Substitution with F gives

greater structural stability. It has a closer coordination than the hydroxyl to the

nearest calcium. Flouride treatment increases bone formation and comprehensive

strength of osteoporetic tissues. The leaching of F from the ceramic might cause

them to become incorporated into the surrounding bone there by strengthening it.

The P- may also stabilize the calcium phosphate phase present at the interfacial

bonding. Other ionic substituents are C03

2" and HP04

2-.

The CaP phases precipitated from aqueous solutions mainly include

dicalcium phosphate [CaHP04.2H20, DCPD], octacalcium phosphate [Ca8

(HP04)2(P04)4.5H20, OCP] and HA. HA is considered as the most

thermodynamically stable phase in physiological environment. OCP and DCPD

have been regarded as precursors of HA or the metastable phases of Ca-P because

that they are kinetically favourable [97-102]. OCP is one of the precursors during

the biomineralization process [103]. Tricalcium phosphate [Ca3(P04)2] is very

similar to HA. The concurrent existence of both HA and J3-TCP forms biphasic

calcium phosphate that combine the excellent bioactivity of HA with the good

resorbability of P-TCP, and thus they are used as bone replacement materials [ 104-

106].

One of the essential criteria in the apatite based materials for medical

application is their aptitude for maintaining thermal stability. Some of their critical

applications are 1) porous or granulated materials useful in bone surgery, 2)

additives to organic polymers that improve their biofunctionality and 3) coatings,

28

plasma sprayed or prepared by laser ablation, need treatments at elevated

temperatures. The pure HA have thermal stability up to 1200°C [ 107]. Porous

ceramic have low mechanical strength and is less suitable for loaded conditions.

But they have excellent biocompatibility [108]. HA is favourable for the stability

during the longer period, amorphous CaP is advantageous for only the

osteoconductive property during the initial fixation of porous materials [109]. The

pH of the electrolytic bath influences the deposition of calcium phosphate. At low

pH a calcium deficient apatite is get deposited. The low molar ratio was due to the

substitution of Ir for Ca2+ during long time electrolysis in the bath of pH about 3

[110]. Calcium phosphate ceramics are also deposited from S. B. F.[111].

Dense CaP ceramics are used for the reconstruction of bone defects. But the

main disadvantage of them is their difficulty to be properly delivered and molded

in order to achieve a desirable contour. Also the calcium phosphate ceramics must

have biocompatibility with bone tissue as well as the surrounding soft tissue [112].

Amorphous CaP was precipitated at 20°C from highly super saturated solutions

having compositions of 1.5� Ca/P� 2.0 at pH 11. Higher Ca/P ratio of the starting

solution resulted in faster crystallization to HA. Shorter induction time for such

crystallization is due to the smaller particle sizes of the initially precipitated

amorphous calcium phosphate as well as due to the Ca rich environment [113]. At

higher current densities and after longer time when the Ca/P ratio of the electrolyte

is moved away from the stoichiometric 1.67 ratio the transformation of amorphous

calcium phosphate to HA with needle like morphology takes place [114].

HA has similarity to bone and finds departure from the chemistry of pure

apatite. It is the ideal match to living bone [ 115, 116]. HA is extensively used in

29

the repairing, reconstruction and replacement of damaged parts of the body [ 11 7-

119]. HA in long bones and tooth are oriented in the direction of the c-axis. This

orientation contributes to the chemical stability of HA crystals in vivo [120].

In resorbable prosthesis, the implant is remodeled by osteoclast activity and

is eventually replaced by osteoid. It is replaced by normal functional bone thus

eliminating any long term biocompatibility problems. But during the remodeling

process the load bearing capacity of the prosthesis is significantly weakened and

mechanical failure may result and is therefore used in temporary fixation devices.

Another drawback is the complicated fabrication process and particularly difficult

shaping.

1.5. CORROSION OF METALLIC IMPLANTS

Corrosion is one of the major processes that cause problems when metals

and alloys are used as implants in the body [121]. Corrosion of implants in the

aqueous medium of body fluid takes place via electrochemical reactions [122]. The

electrochemical reactions that occur on the surface of the surgical alloy are

identical to those observed during the exposure of sea water (aerated sodium

chloride). The metallic components of the alloy are oxidized to their ionic forms

and the dissolved oxygen is reduced to hydroxyl ions. During corrosion the total

rate of oxidation and reduction reactions are equal. Corrosion releases significant

concentrations of corrosion products into the solution.

The metals and alloys used as surgical implants achieve passivity by the

presence of a protective surface passive film. This film inhibits corrosion and

keeps current flow and the release of corrosion products at a very low level. i. e. all

30

the implantable materials undergo corrosion at some finite rate due to complex

corrosive environment in the body. The types of corrosion that are pertinent to

currently used alloys are pitting, crevice, galvanic, inter granular, stress-corrosion

cracking, corrosion fatigue and fretting corrosion [123].

a) Pitting corrosion

If the passivation film breaks down, corrosion takes place at the point,

which becomes anodic, whilst the rest of the material becomes cathodic. Pitting is

the accelerated corrosion. It is a severe form of localized corrosion and results in

extensive damage and release of significant amount of metal ions. Pitting refers to

the formation of small cavities/holes at the surface of a material, which is protected

otherwise by the presence of an adherent, tenacious and self-healing thin passive

film. The formation of such pits is attributed to the interaction of certain aggressive

ions within the film at locations where it is defective or weak in nature. The pits

may be visible to the naked eye in some cases but in general they are invisible, and

dangerous to the extent they can allow the formation of stress corrosion cracking

(SCC) or fatigue cracks, which can catastrophically fail the components in service.

The importance of pitting significantly depends on the nature of the surface layer

or the film that has formed on the surface due to the interaction of the material with

the environment. Thus a state of passivity is forced into the material, which

safeguards the material from general corrosion by slowing down the dissolution

process at the surface.

In implants, pitting occurs most often on the undesirable screw heads. This

form of attack occurs more frequently in media containing chloride ions [ 124]. It is

well established that the resistance to pitting in saline environment can be

31

increased by molybdenum addition and keeping the inclusion contents to a

minimum level.

b) Crevice corrosion

Crevice corrosion is a form of corrosion, related to structural factors. It

occurs when a metal surface is partially shielded from the environment. It is

usually encountered beneath the screw head that holds the plate or similar locations

such as the intersection of the components of two pieces, hip nail etc. The basic

requirement for the occurrence of this process is the presence of a crevice, a

narrow deep crack: either an interface between parts of a device, such as between

plate and screw head, or defects such as fatigue crack. Type 316L SS is highly

susceptible to crevice corrosion attack as compared to other commonly used

metallic implant material [ 125]. The occurrence of corrosion on the bone plate and

screws made of SS especially in the area of contact between screw heads and

counter sink hole is a common feature. The presence of crevice corrosion in the

counter sink portion of the bone plate can also induce crack propagation: though

such type of crevice induced failures have been seldom identified. Crevice

corrosion problem can often be eliminated by appropriate design of device and

proper choice of material.

c) Galvanic corrosion

Galvanic or bi-metallic corrosion takes place when two different metals are

m physical contact in an ionic conducting fluid medium such as serum or

interstitial fluid. The differential composition or process variables of a plate and

the adjoining screws is responsible for the set-up of a galvanic couple, which

results in galvanic corrosion. Galvanic corrosion depends on a large number of

32

complicating factors such as the relative areas of electronic and ionic contact, and the actual metal pair involved. However, it is safe to assume that some galvanic corrosion will occur if a bone plate and bone screw is made of dissimilar metals or alloys. Corrosion is likely to occur between the plate and bottom side of the screw holes. d) Corrosion fatigue

Corrosion fatigue is a fracture failure of metals that occurs because of the combined interaction of electrochemical reactions and cyclic loading. Corrosion fatigue resistance is an important factor of consideration for load-bearing surgical implant metals or for more metals used in cyclic motion applications. Normally, a failure may not occur, but cracks can initiate from hidden imperfections, surface damage, chemical attack and other causes. The corrosive environment may result in local corrosive attack that accentuates the effect of the various imperfections. The corrosive attack will be influenced by solution type, solution pH, oxygen content and temperature. The body fluid environment may decrease the fatigue strength of the implant. Fatigue striations are observed on the fractured surface of the device with coloured "beach marks" are indicative of corrosion fatigue. The presence of corrosion pits could induce the fatigue to develop [126]. Failures of mechanical origin in orthopaedic implants are most commonly due to fatigue or environmentally assisted fatigue. In some instances, however, the mechanism responsible for crack initiation and crack propagation may be different [127]. It has been suggested that cracks could be initiated by fretting and propagated by stress­corrosion cracking (SCC). It is also possible that cracks could initiate because of corrosion phenomena and propagated by a fatigue mechanism.

33

e) Fretting Corrosion

Fretting corrosion occurs when two opposing surfaces such as bone plates

and the screw heads of the prosthetic devices rub each other continuously in an

oscillating fashion in the body environment. It is the result of small relative

movements between the contacting surfaces in a corrosive medium. Even in the

absence of corrosive medium, fretting can occur. Clinical significance of fretting

attack lies in its intensity that may give rise to a large amount of corrosion products

in adjacent tissues or it may be a major factor in crack initiation and fracture failure

of an implant [128]. Fretting corrosion at counter sinks in plates and at hip nails

may initiate corrosion fatigue, which occurs through hole. Quantification of weight

loss of the implant due to fretting corrosion has been studied and was found to be

directly proportional to the load transmitted across the surfaces, the number of

cycle fretted and the amplitude of stresses. The weight loss has been reported to be

inversely proportional to the hardness of the material and the frequency of stroke.

1.5.1. Mechanism of corrosion

Reactions of metals with aqueous environments are electrochemical in

nature involving movement of electrons to the cathode. For implanted metals in

aqueous environment with dissolved oxygen (Fig. 1.4) the primary anodic and

cathodic reactions are represented respectively by

M

Anode :rvr++

i

\ �':\I.\ .. .

···\·._- -: : : : : e· .. : :: .-: _.·.... e- ..

Cathode

20H-

j

Metal

34

Fig. 1.4. Schematic illustration of electrochemical set up between

anodic and cathodic sites on an iron surface undergoing corrosion

The crevices between components, wounds etc., can have extremely low

concentration of oxygen leading to cathodic reactions of water.

Thus most corrosion in metals occurs through the oxidation process at

anode. The metals with positive potential are the noble metals which are least

reactive (cathodic). The corrosion rate is directly related to the current flow

between the anode and cathode. The electrode potential with respect to solution is

a measure of Gibb's free energy of the reaction.

�G =-nEF

where E is reaction potential and F is Faraday's constant (amount of

electricity associated with the flow of electrons, 96,487 C/mol. Equivalent). In the

galvanic corrosion the transfer of electrons occurs due to the difference in

composition, energy level of electrolytic environment. Table 1.4. lists out the

galvanic series for some implant metals[ 129].

Table 1.4. The galvanic series for some implant metals

(-) Cathodic Gold

Noble, least reactive Graphite

Silver

(+) anodic

316L SS (passive)

304 SS (passive)

Titanium

316L SS (non psssive)

Active, most reactive Aluminium

35

The variation in oxygen concentration over the surface in the environment

can induce electrochemical cell. The sites with low oxygen concentration become

anodes and corrosion takes place. The electrode potential of a metal can be altered

by its thermo mechanical state. The potential energy at grain boundaries is higher

than that in the middle of a grain causing the grain boundaries or second phase to

be anodic resulting in corrosion. The corrosion can be altered in the presence of

state or dynamic stress. Any region of distortion or stress becomes anodic with

respect to unstressed region which has a lower energy level leading to stress or

fatigue corrosion.

Pitting corrosion occurs when the anodic site becomes fixed on a small part

of the surface due to a break in the passive layer, irregularities in the surface

caused due to variations in the metal itself, incomplete coatings, scaling deposits

that build up at selected areas, etc. The cathodic area can be anywhere outside the

pit. This results in a high corrosion current density at the base of pits. Movement of

36

metal ions or Ir ions from the bottom of the pit is restricted by the film covering

the top of the pit [130]. The depletion of oxygen creates a difference in

electrochemical potential between the pit and the surrounding metal. The formation

and breakdown of passivity and the mechanism of pitting attack are shown in Fig.

1.5. Once the pit is initiated, the metal ions form precipitates at the top of the pit

and often form a film covering the pit. The film resists the entry of the solution and

oxygen into the pit.

Anodic reaction: Metal- e· ---+ M Cathodic reaction: % O::i + H::iO + 2 e· ___.,. 20ff

o= H::iO Chloride � _ Breakdown

Cl

Passivity

(b)

Reference half-cell

Reference half-cell

Fig.1.5. Two principles of break down of passivity: (a) Metal undergoes

general corrosion with film dissolution; (b) metal undergoes pitting corrosion.

In all highly alloyed metals in the body environment, galvanic corrosion

can occur but to a lesser extent than in the iron based alloys.

The metallurgical analysis of failed SS implants indicates a variety of

deficiency factors is responsible for the failure. These include the deficiency of

37

Mo, the use of sensitized steel, the inadvertent use of mixed metal and use of

mutually incompatible components.

1.5.2. Prevention of corrosion

In young patients with higher life expectancy new questions in relation to

long term use of implants results. Local and systemic tissue reactions to the

implants are of interest. Tissue reactions will lead to alien element induced peri­

prosthetic inflammation, osteolysis, and activation of the reticuloendothelium

system as well as a potential carcinogenic effect triggered by degradation products

of implanted prosthetic materials [131-133].

The corrosion of metallic implants can affect the surrounding tissue in three

ways. 1) electrical current may affect the behaviour of the cells, 2) the corrosion

process may alter the chemical environment, 3) the metallic ions may affect

cellular metabolism. Hence the prevention of corrosion of metallic materials is

very much important. The metals normally react with oxygen in the air and forms a

passive oxide layer. Passivation of metals i.e., the formation of surface oxide film

improves the corrosion resistance. SS forms chromium oxide surface film. Also

the Ti metal forms Ti02 passive film. The Ti02 is biocompatible and it can

enhance apatite formation. The corrosion of metallic substrates is better prevented

by providing a bioactive layer on metallic substrate. Self healing films can reform

after being damaged. This results in high interfacial stiffness and strength with a

direct bond between bone and bioactive material.

1.5.3. Coating of bioactive materials on implants

The corrosion of metallic implants can be best prevented by providing a

coating on the surface of the metal. Since the metallic implants are continuously in

38

direct contact with the physiological solution, the coating must be compatible to

the aggressive biological environment. To overcome loosening problems implants

are coated with bioactive ceramics. Biologically compatible and active materials

are usually coated on the metallic substrates. Bioactive materials chemically bond

to tissues through a layer of biologically active apatite that forms in vivo [12].

Bioceramic coating on metallic substrates have been widely used in medicine and

dentistry to combine the excellent mechanical properties of metal alloys with the

bioactive properties of bioceramics [134]. High interfacial stiffness and strength

with a direct bond with bone and bioactive material is achieved through the fibrous

interface between bone and non bioactive materials [135, 136]. Biologically active

coating is used to modify metal implants and to endow metal materials with bio

activity and to improve interfacial strength between the implant and the host. An

ability to control cell response at the biological-material interface is required to

manufacture a new generation of biomaterial that are now in demand.

Calcium phosphate ceramics are widely used as bone substitutes m

dentistry, orthopaedics and reconstructive surgery. It is biocompatible and

osteoconductive. These ceramics are only available as pre-fabricated blocks or

granules. Pre-fabricated blocks are difficult to shape, resulting in poor filling of the

bone defect, while granules do not provide the dimensional stability and can easily

migrate into the surrounding tissue. Solution for this problem is that calcium

phosphate cement can be shaped into the defect dimension and harden in situ [ 13 7-

140]. The favourable material for coating on metallic substrate is natural bone

mineral, hydroxyapatite (HA). Hydroxyapatite (HA) is a particularly attractive

material for human tissue implantation. The load bearing implants cannot be

39

entirely made of HA because it is a brittle ceramic. The HA coating provide better

tolerance of the coating in the body condition. A HA coating is preferable for the

stability during the longer period, and an amorphous calcium phosphate (ACP)

coating is advantageous for only the osteoconductive property in the initial fixation

ofporous materials [141]. HA coated bio inert implants can provide well behaved

bone substitutes in bioactive and mechanical properties. HA is osteoconductive.

The osteoconductivity of HA is due to the RGD peptide domain dependent

spreading of osteoblasts [142]. HA coating induces reliable fixation relatively early

by means of chemical bonding [143-146].

Surface oxide properties have great importance in establishing successful

osteointegration of Ti implants. The oxide thickness, micro pore structure and the

crystallinity significantly influence the bone tissue response [147]. Hence the

surface pre-treatment of Ti is significant in improving the adherence of HA

coating. Addition of Ti02 in the HA matrix can enhance the mechanical properties

of the composite [148]. Commercially pure Ti plates were heated in air and then

immersed in a super saturated calcium phosphate solution and the calcium and

phosphate ions precipitated on to the Ti02 film. Heat oxidation enhanced the

reactivity by increasing the surface roughness, surface energy and amount of

hydroxyl groups [149]. All apatite inducers have a great number of acidic hydroxyl

groups on their surfaces that included SiOH, TiOH and COOH. A material could

induce apatite formation when there are a large number of acidic OH groups

present on their surfaces [ 150]. Alkaline treatment of Ti before HA coating is

essential to induce hydroxyl groups on the surface. Bioglass has high bioactivity

due to rapid formation of hydroxycarbonate apatite on the biomaterials surface in

40

S. B. F. and the different behaviours of ion diffusion and bio-mineralisation in S.

B. F. solution due to their different composition [151]. Hence bioglass coatings are

used on the surface of metallic implants. HA coatings on alumina are developed

under hydrothermal condition from an aqueous solution of calcium nitrate and

ammonium dihydrogen phosphate containing EDT A sodium salt and hydrogen

peroxide [84]. The incorporation of Ti02 component into a PTMO-CaO-Si02

hybrid results an increase in the apatite forming ability in S.B.F. [152]. HA/

bioglass enhances direct bonding between the coating and the bone. Colloidal

silica deposited on glass substrate is a simple model substrate for osteointegration

[153]. CaP coating on Ti6Al4V with an intermediate silica barrier layer was

deposited by CVD method using a metal and organic precursor [154].

During the formation of bioactive layer of HA at body temperature

antibiotic and growth factors that are stable in acid medium is introduced. The

incorporation of chitosan into HA will increase the biocompatibility of the

electrolytic coating while maintaining its original strength [155]. The apatite layer

coating on Ti alloys in situ may improve bone-to-implant bonding make it as a

promising coating material [156]. In addition an octacalcium phosphate (OCP,

CasH2(P04)6.5H20) coating was recommended as a biomimetic coating for use in

orthopaedic surgery because OCP was known to be one of the precursors during

the bone mineralization process [110]. With chemical treatments prior to

deposition the coating sometimes get more uniform and well adhered [157]. Hence

the metallic substrates are subjected to oxidation treatments. HA coating can

provide reliable early fixation by chemical bonding and in combination with

porous metals allows rapid and strong fixation with bone. HA/Zr02 coating

41

could be deposited on Ti by plasma spray method [92]. Flourapatite based glass

ceramic coating on alumina provide good sites for the nucleation and growth of an

apatite layer equivalent to the mineral component of bone and binds preferentially

plasma fibronectin which is well known to enhance cell adhesion and spreading

[85]. Glass ceramic coating on 13-Ti alloy give higher bond strength than pure Ti or

Ti6Al4V [158]. Bioactive glass ceramic coatings on alumina give it the best

osteoblast like cell behaviour [159].

Techniques of HA coating are plasma spraymg [160, 161],

electrodeposition [162], sol-gel method [163], laser ablation [164], ion beam

sputtering [165], solution combustion synthesis [166] and biomimetic coating

[167-172]. Plasma spraying is the best method to produce HA coating, but the

chemical composition and crystallographic phase of the coating cannot be

controlled with this method [173]. Thermal plasma spraying provides high bond

strength between the substrate and the coating. Coating by electrophoresis appears

to be an attractive process to coat metallic implants with an osteoconductive

material like HA [174]. Plasma spraying cannot produce a uniform coating on

devices with complex shapes. Therefore to improve these disadvantages

electrodeposition method is adopted. It is a quick and uniform coating method. The

method gives uniform coating on substrates with complex shapes at low

temperature and the control of film thickness and composition could be achieved

by this method [175-179].

Biomimetic deposition induces calcium phosphate coating on Ti surface in

aqueous solution at low processing temperature [55,180-183]. Biomimetic

deposition is the method in which materials are deposited from a biological

42

solution in a controlled condition. Biomimetic deposition was carried out in

simulated body fluid (S. B. F.) that has composition similar to human blood

plasma. The mechanism suggested for apatite deposition in S. B. F. is the release

of titaniumhydroxide and OH ions from the Ti surface and their accumulation

inside the confined space between the two contact surfaces [184]. HA precipitation

exhibit higher thermodynamic driving force than OCP and DCPD in S. B. F. The

HA nucleation rate is affected by pH value. High pH environment is favourable for

HA nucleation and the HA nucleation rate approaches the nucleation rates of OCP

when the pH value approaches 10. DCPD does not have a thermodynamic driving

force of precipitation in S. 8. F., even though it has kinetic advantages. DCPD

precipitation is possible when the concentration of c/+ and PO/- ions increase to

a higher than normal level in S. B. F. [185].

Of the various methods of preventing the corrosion of metallic implants the

HA coating is the most suitable one. For better adherent HA coating pre-treatment

of the substrates is necessary.

1.6. CURRENT RESEARCH TRENDS

Corrosion is one of the major problems resulting in the failure of metallic

implants. Hence the prevention of corrosion of metallic implants gets very high

significance. A lot of research is going on to develop methods to prevent the

corrosion of metallic implants. The nature of passive films formed is an important

criterion in preventing the corrosion of implants. Various new corrosion resistant

materials are developed. A new class of super austenitic SS, which have a higher

43

percentage of Cr and Mo content, strengthens the passivity and enhances the

corrosion resistance. Nickel free SS is developed to eliminate the toxicity of nickel.

Vanadium is harmful to human body. Hence vanadium free Ti alloy-Ti6Al7Nb­

having improved wear and corrosion resistance is produced. The elastic modulus

of bone is 30 GPa whereas that of Ti alloys is 120 0Pa. Hence new class of near

beta-Ti alloys is developed and it is non toxic also. Controlling the microstructure

by heat treatment is an effective way to improve the properties of alloys especially

corrosion resistance properties.

In biomimetic method chemically treated Ti is immersed m a

supersaturated calcification solution containing proteins to directly form

carbonated HA containing proteins. Also biomimetic method of coating is

modified with an approach to reduce the time period required for apatite formation.

Highly concentrated S. B. F. was developed for fast biomimetic deposition.

The organic-inorganic hybrid materials [PTMO-CaO-Si02] may be useful

as a new kind of bioactive bone repairing material [ 152]. The bioactive glass

coating is an emerging field in orthopaedic implants. Bioacive glass coatings on

metals enhance the bicompatibilty of metallic substrates. Also new kinds of

ceramic materials were developed having high biocompatibility.

Ultrafine/nano phase metals come into the growing list of materials

enhanced bone cell function pertinent to successful orthopaedic implant

applications were observed [186]. Nano HA has no cytotoxic effect and is

convinced by in vivo experiments [187]. The HA around bone is found to have

nano size. Nano· structured products were developed by innovative synthetic

technique capable of yielding powders with higher reactivity and bio activity

44

[188]. Nano grade needle like crystals of CaP having composition and morphology

similar to apatite crystals in bone was developed and forms bone like implant

[189].

The problem of adherence of the HA coating to the metallic substrates is

improved by providing pre-treatments to the substrate surface and by modifying

the deposition process. Ion implantation is another method to improve the

biocompatibility of the metallic substrates.

Orthopaedic implants is now an important research area having very high

potential to get modified and to develop the most biocompatible and economic

coating.

1.7. THE PRESENT STUDY

The present study deals with the development of HA coating having high

stability. A uniform HA coating was developed on Ti substrate by modification of

electrodeposition process. The electrodeposition process was modified by adjusting

the throwing power of the electrolyte. Throwing power represents the ability of the

plating bath to give an even deposit. In the present study the throwing power was

controlled by varying the current density and inter electrode distance. By adjusting

the inter electrode distance and current density to the optimum value the desired

phase HA is obtained.

The microstructure of the deposited HA matrix was modified by

incorporating clay into the HA coating. The clay incorporated HA coating behaved

similar to the biological apatite since it contains trace elements included within the

crystal structure of the clay. The clay was incorporated to get the HA coating

45

having the similar property of biological HA. Also the sodium, magnesium ions

incorporated in the coating matrix had certain biological functions. The silica and

hydroxyl content of the coating also enhances the further apatite deposition from S.

B. F.

The adherence of the HA coating was improved by providing an adherent

interlayer. HA/Ni biomedical coating is developed by electrolessdeposition

technique [190]. In the present study nano HA incorporated electroless Ni-P

coating was developed for further HA coating. The nano HA was incorporated into

the electroless bath and the coating obtained was very much adherent to the SS

substrate. It is found to be corrosion resistant in physiological solution.

Zinc phosphate cements have been used in dentistry for many years and as

early as 1879 as joint material between tooth root and crown [191]. Zinc releasing

CAP favour human bone marrow cell culture and bone formation [192]. Zinc

releasing CaP promoted bone formation when implanted in rabbit [193]. In the

present study a zinc phosphate inter layered HA coating is developed to enhance

the adherence of the HA coating. The SS substrate is hot-dip galvanized prior to

conversion coating. Further HA coating was developed on the conversion coated

zinc phosphate.

The developed coatings were evaluated in detail by physiochemical,

morphological and electrochemical evaluation. The developed coatings were

evaluated and are found to be corrosion resistant in nature. The biomimetic

evaluation of the developed coatings in S. B. F. revealed the biocompatibility of

the developed coatings.The present study gives a better insight into the

development of HA coating having superior performance.

46

1.9. REFERENCES

[1] R. J. Gillespie, D. A. Hurnpreys, N. C. Baird, E. A. Robinson, "Chemistry",

2nd ed., (Prentice Hall, Upper Saddle River, NJ, 1989).

[2] M. J. H. Smith, L. V. Interrante, "Introductory Terms and Concepts" in

"Chemistry of Advanced Materials an Overview" edited by L. V.

Interrante, (Wiley-VCH. Inc., 1998).

[3] A. Tampieri, G. Celotti, E. Landi, Anal. Bioanal. Chem. 381 (2005) 568.

[4] H. B. Pogge, Ed., "Electronic Materials Chemistry", (Marcel Dekker, New

York, 1996).

[5] H. W. Salzberg, "From Caveman to Chemist", (American Chemical

Society, Washington DC, 1995).

[6] "Materials and Man's Needs", Summary report of the Committee on the

Survey of Materials Science and Engineering, (National Academy of

Sciences, Washington DC, 1974).

[7] D. F. Williams, Definition ofBiomaterials, (Elsevier, Amsterdam, 1987).

[8] C. J. Kirpatrick, F. Bittinger, M. Wagner, H. Kohler, T. G. van Kooten, C.

L. Klein, M. Otto, Pro. Instn. Mech. Engrs. 212 Part H (1998) 75.

[9] J. B. Park, "Biomaterials Science and Engineering," (Plenum Press, New

York, 1984).

[10] Spotnitz, "Circulatory Assist Devices, in Handbook of Bioengineering"

(McGraw-Hill, New York, 1987).

[11] P. S. Christel, Clin. Orthop. Relat. Res. 282 (1992) 10.

[12] L. L. Hench, J. Am. Ceram. Soc. 74 (1991) 1487.

47

[13] F. Hulbert, in "An introduction to Bioceramics", (World Scientific,

Singapore, 1993).

[14] R. Cordingley, L. Kohen, B. Ben-Nissan, G. Pezzotti, J. Australasian

Ceram. Soc. 39 (2003) 20.

[15] W. Suchanek, M. Yoshimura, Pro. J. Mater. Res. 1 (1998) 94.

[16] M. Jarcho, Clin. Orthop. Relat. Res. 157 (1981) 259.

[17] H. M. T. U. Herath, L. DiSilvio, J. R. G. Evans, J. Mater. Sci: Mater. Med.

16 (2005) 1061.

[18] G. Willman, in "Bioceramics in Joint Replacement: State-of-the-art and

Future Options". Ceramic Forum International 79 (2002) E27.

[19] H. Y. Yasuda, S. Mahara, T. Nishiyama, N. Terashita, Y. Umakoshi,

"Tissue Engineering for Therapeutic Use 6", International Congress

Science Series, 1243 (2002) 155.

[20] D. M. Magarita, L. Peter, S. Istvan, Adv. Sci. Technol. 12 (1995) 321.

[21] S. H. Rhee, J. Y. Choi, H. M. Kim, Biomaterials 23 (2003) 4915.

[22] S. Rhee, J. Biomed. Mater. Res. Part A 67 (4) (2003) 1131.

[23] H. Takadama, M. Hashimoto, Y. Takaigawa, K. Jono, M. Mizuno, Y.

Yasutomi, T. Kokubo, N. Miyata, Zairyo 52 (2003) 581.

[24] 0. Wichterle, D. Lim, Nature 185 (1960) 117.

[25] J. Kopecek, J. Yang, Polym. Int. 56 (2007) 1078.

[26] A. Oyane, M. Kawashita, T. Kokubo, M. Minoda, T. Miyamoto, T.

Nakamura, J. Chem. Soc. of Japan 110 (2002) 248.

[27] J. E. Davies, Biomaterials 28 (2007) 5058.

48

[28] C. Rosse, P. G. Rosse, in "Hollinshead's Text book of Anatomy,"

(Lippincott, Williamson and Wilkins, London, UK, 1997).

[29] T. Kirkwood, in "Times of Our Lives," (Widenfeld and Nicholson,1999).

[30] L. L. Hench, MRS Bull. 24 (1999) 13.

[31] B. Gullberg, 0. Johnell, J. A. Kanis, Osteoporosis Int. 7 (1997) 407.

[32] M. J. Parker, Z. Gerontol. Geriat. 34 (2001) 74.

[33] E. W. Abel, in "Hip Arthroplasty," Distance learning Section, Dept. of

Orthopaedic and Trauma Surgery, University of Dundee (1997).

[34] M. Niinomi, in "Structural Biomaterials for the 21st Century", edited by M.

Niinomi, T. Okabe, E. M. Taleff, D. R. Lesuer H. E. Lippard, (TMS,

2001).

[35] B. 0. Ratner, in "Titanium in Medicine", edited by D. M. Brunette, P.

Tengvall, M. Textor, P. Thompsen, (Springer, Berlin, 2001).

[36] Y. Z. Huang, D. J. Blackwood, Electrochim. Acta 51 (2005) 1099.

[37] J.B. Cotton, Chem. Eng. Prog. 66 (1970) 57.

[38] D. Schlain, J. S. Smetko, J. Electrochem. Soc. 99 ( 1952) 417.

[39] V. V. Andreava, Glukova, J. Appl. Chem. 11 (1961) 390.

[40] T. Albrektsson, P. I. Branemark, H. A. Hansson, B. Kassemo, K. Larsson,

I. Lundstrom, D. H. Mcqueen, R. Skalak, Ann. Biomed. Eng. 11 (1983) 1.

[41] D. F. Williams, in "Biocompatibility of Clinic Implant Materials", edited

by D. F. Williams, (CRC Press, Boca Raton, Florida, USA, 1981).

[42] D. F. Williams, in "Titanium in Medicine" edited by D. M. Brunette, P.

Tengvall; M. Textor, P. Thopsen, (Springer, Berlin, 2001).

49

(43] L. L. Hench, in "Bioceramics-7", edited by 0. H. Anderson, A. Yli Urpo,

(Butterworth-Heinemann, Oxford, England, 1994).

(44] R. Z. Leggeros, R. G. Craig, J. Bone Miner. Res. 8 (1993) S 583.

(45] S. G. Steinemann, lnjuri 27 (1996) SC 16.

(46] K. Suzuki, K. Aoki, K. Ohya, Bone 21 (1997) 507.

(47] J. J. Jacobs, J. L. Gilbert, R. M. Urban, J. Bone Joint Surg. Am. 80 (1998)

268.

(48] T. Sawase, K. Hai, K. Yoshida, K. Baba, R. Hatada, M. Atsuta, J. Dent.

Res. 26 (1998) 116.

[49] B. Kasemo, J. Prosthetic Dent. 49 (1983) 832.

(50] P. Ducheyne, P. Bianco, S. Radin, E. Schepers, In "Bone-bonding

Biomaterials", edited by P. Ducheyne, T. Kokubo, C. A. van Blitterswijk,

(Leiderdorp, Reed Healthcare Communications, The Netherlands, 1992).

[51] C. M. Lin, S. K. Yen, J. Meta Stable and Nano Crystalline Materials 23

(2005) 121.

(52] T. Kokubo, H. M. Kim, H. Takadama, M. Uchida, S. Nishiguchi, T.

Nakamura, Metal. Mater. Res. Symp. Proc. (Mineralization in Natural and

Synthetic Biomaterials) 599 (2000) 129.

(53] M. Uchida, H. M. Kim, T. Kokubo, T. Nakamura, Sixth World Biomater.

(2000) 130.

[54] K. S. Hwang, Y. H. Lee, B. A. Kang, S. B. Kim, J. S. Oh, J. Mater. Sci.:

Mater. Med. 14 (2003) 521.

[55] D. M. Brunette, P. Tengwall, M. Textor, P. Thomsen, "Titanium m

Medicine," (Springer Verlag, 2001).

50

[56] B. Liang, S. Fujibayashi, M. Neo, J. Tamuraa, H. M. Kim; M. Uchida, T.

Kokubo, T. Nakamura, Biomaterials 24 (2003) 4959.

[57] P. van Landuyt, J. M. Streydio, F. Delannay, E. Munting, Clin. Mater. 17

(J 994) 29.

[58] B. C. Yang, 1- 'f. Uchida, H. M. Kim, X. Zhang, T. Kokubo, Biomaterials 25

(2004) 1003.

[59] A. Osaka, S. Hayakawa, K. Keiko, K. Tsuru, Mem. Fae. Eng., Okayama

Univ. 33 (1999) 79.

[60] K. Tsuru, S. Takemoto, S. Hayakawa, A. Osaka, Mater. Res. Symp. Proc.

599 (2000) 141.

[61] Q. Zhang, Y. Leng, Biomaterials 26 (2005) 3853.

[62] S. Nishiguchi, T. Nakamura, H. K. Kim, F. Miyaji, T. Kokubo,

Biomaterials 20 (1999) 491.

[63] H. B. Wen, J. G. Wolke, J. R. Wijn, Q. Liu, F. Z. Cui, K. de Groot,

Biomaterials 18 (1997) 1471.

[64] H. M. Kim, F. Miyaji, T. Kokubo, T. Nakamura, J. Biomed. Mater. Res. 32

(1996) 409.

[65] H. M. Kim, F. Miyaji, T. Kokubo, S. Nishiguchi, T. Nakamura, J. Biomed.

Mater. Res. 45 (1999) 100.

[66] T. Kokubo, F. Miyaji, H. K. Kim, T. Nakamura, J. Am. Ceram. Soc. 79

(1996) 1127.

[67] H. M. Kim, F. Miyaji, T. Kokubo, T. Nakamura, J. Mater. Sci.: Mater.

Med. 8 (1997) 341.

51

[ 68] P. Habibovik, F. Barrere, C. A. van Blitterswijk, K. de Groot, P. Layrolle,

J. Am. Ceram. Soc. 85 (2002) 517.

[69] M. Sato, E. B. Slamovich, T. J. Webster, "Ceramic Transactions" 159

(Ceramics, Nanomaterials and Nanotechnology III, 2005), 231.

[70] R. J. Solar, "Corrosion Resistance of Titanium Surgical Implant Alloys: A

Review", in: S. C. Syrett, A. Acharya, editors. "Corrosion and Degradation

of Implant Materials". (Philadelphia: ASTM Special Technical

Publications. STP 684, 1979) 259.

[71] M. Ask, J. Lausmaa, B. Kasemo, Appl. Surf. Sci. 35 (1989) 283.

[72] C. Johansson, J. Lausmaa, M. Ask, H. A. Hansson, T. Albrektsson, J.

Biomed. Eng. 11 ( 1989) 3.

[73] J. Lausmaa, B. Kasemo, H. Mattsson, H. Odelius, Appl. Surf. Sci. 45

(1990) 189.

[74] R. L. Reis, 0. C. Paiva, J. D. Santos, Bioceramics 7 (1994) 125.

[75] H. Oonishi, E. Tsuji, T. Sujimoto, "Comparison of Bone Ingrowth into

Ti6Al4V Beads Coated and Uncoated with Hydroxyapatite Orthopaedic

Ceramic Implants", Japan, 6 (1986) 87.

[76] J. W. Oh, H. C. Choe, Y. M. Ko, Han'guk Pyomyon Konghak Hoechi,

Korean Institute of Surface Engineering, 37 (2004) 249.

[77] C. C. Shih, C. M. Shih, Y. Y. Su, L. H. J. Su, M. S. Chang, S. J. Lin,

Corros. Sci. 46 (2004) 427.

[78] T. M. Sridhar, U. Kamachi Mudali, Transactions of the Indian Institute of

Metals 56 '(2003) 221.

52

[79] S. Kannan, A. Balamurgan, S. Rajeswari, Electrochim Acta 50 (2005)

2065.

[80] I. Gurappa, Surf. Coat. Technol. 161 (2002) 70.

[81] X. Miao, Mater. Lett. 57 (2003) 1848.

[82] · J. X. Liu, Y. D. Li, L. Yu, Dalian Qinggongye Xueyuan Xuebao 23 (2004)

160.

[83] T. Hanawa, in "Recent Research Developments in Biomaterials" edited by

Y. Ikada (Research Sympost, Korea 2002) p.12.

[84] J. S. Ha, M. N. Rahaman, "Formation of Hydroxyapatite Coatings on

Aluminia Under Hydrothermal Conditions". Ceram. Eng. & Sci. Proc.

2003, 24 (3, 27th International Cocoa Beach Conference on Advanced

Ceramics and Composites, Part A, ACS, 2003) 239.

[85] M. Bosetti, E. Verne, C. V. Brovarone, C. Moisescu, M. Sabbatini, M.

Cannas, J. Biomed. Mater. Res. 66A (2003) 615.

[86] H. Noda, K. Muramoto, H. Kim, J. Mater. Sci. 38 (2003) 2043.

[87] F. Barrere, A. Lebugle, C. A. van Blitterswijk, K. de Groot, P. Layrolle, C.

Rey, J. Mater. Sci: Mater. Med. 14 (2003) 419.

[88] T. Wei, A. Ruys, B. Milthorpe, Key Eng. Mater. 240-242 (2003) 591.

[89] A. RapaczK.mita, A. Slosarczyk, Z. Paszkiewicz, Ceramics International 31

(2005) 567.

[90] B. Vaidyanathan, K. A. Kumar, J. N. Shetty, K. J. Rao, Bull. Mater. Sci. 20

(1997) 135-140.

[91] C. V. Brovarone, F. Smeacetto, E. Verne, Ceram. Eng. & Sci. Proc. 23

(2002) 839.

53

[92] N. Miyata, K. Fuke, Q. Chen, M. Kawashita, T. Kokubo, T. Nakamura,

Biomaterials 23 (2002) 3033.

[93] D. Fan, Z. Xu, S. Li, Y. Yan, Journal of Wuhan University of Technology,

Materials Science Edition 17 (2002) 44.

[94] L. L. Hench, T. K. Greenlee, W. C. Allen, An Investigation of Bonding

Mechanisms at the Interface of a Prosthetic Material-Report#! US Army

Medical R&D Contract no DADA 17-70-C-001 August 1970.

[95] T. Kasuga, Y. Hosoi, M. Nogami, M. Niinomi, Phosphorous Res. Bull.12

(2001) 39.

[96] K. Zhang, L. F. Francis, Y. Han, A. Stein, J. Am. Ceram. Soc. 88 (2005)

587.

[97] R. Z. Leggeros, Prog. Cryst. Growth Charact. 4 (1981) 1.

[98] P. G. Koutsoukos, Z. Amjad, M. B. Tomson, G. H. Nancollas, J. Am.

Chem. Soc. 102 (1980) 1553.

[99] P. G. Koutsoukos, G. H. Nancollas, J. Cryst. Growth 53 (1981) 10.

[100] J.C. Heughebaert, G. H. Nancollas, J. Phys. Chem. 88 (1984) 2478.

[101] J.C. Heughebaert, G. H. Nancollas, Colloid Surf. 9 (1984) 89.

[102] Y. Y. Hiroyuki, M. Shigeo, N. Tsuyoshi. M. Shigeo, N. Tsuyoshi, T.

Naohiro, U. Yukichi, International Congress Series 2002, 1243 (Tissue

Engineering for Therapeutic Use 6), 155.

[103] 0. Gauthier, J.M. Bouler, E. Aguado, R. Z. Legeros, P. Pilet, G. Daculsi, J.

Mater. Sci.: Mater. Med. 10 (1999) 199.

[104] R. Z. Leggeros, S. Lin, R. Rohanizadeh, D. Mijares, J. P. G. Leggeros, J.

Mater. Sci.: Mater. Med. 14 (2000) 201.

54

[105] S. Kannan, J. H. G. Rocha, J. M. U. Ventura, A. F. Lemos, J. M. F.

Ferreira, Chem. Mater. 53 (2005) 1249.

[106] S. Kannan, J.M. F. Ferreira, Chem. Mater. 18 (2006) 198.

[107] D. P. Link, J. van den Dolder, W.J. F. M. Jurgons, J. G. C. Wolke, J. A.

Jansen, Biomaterials 27 (2006) 4941.

[108] M. Nagano, T. Nakamura, T. Kokubo, M. Tanahashi, M. Ogawa,

Biomaterials 17 (1996) 1771.

[109] F. Barrere, P. Layrolle, C. A. van Blitterswijk, K. de Groot, J. Mater Sci.:

Mater. Med. 12 (2001) 529.

[110] M. Tanaka, Y. K. N. Nauri, K. Kuwahara, Journal of the Society of lnorg.

Materials, Japan 10 (2003) 316.

[111] X. Lu, Y. Leng, Biomaterials 26 (2005) 1097.

[112] J. A. Jansen, J. G. C. Wolke, E. M. Ooms, Materials Science Forum (Pt. 4,

THERMEC') 426-432 (2003) 3085.

[113] S. Kim, H. S. Ryu, H. S. Jung, K. S. Hong, Metals and Materials

International 10 (2004) 1 71.

[114] A. Sewing, M. Lakatos, D. Scharnweber, S. Roessler, R. Born, M. Dard, H.

Worch, Key Eng. Materials, 254-256 (2004) 419.

[115] J. C. Eliott, "Structure and Chemistry of the Apatites and Other Calcium

Orthophosphates", (Elsevier Amsterdam, 1994).

[116] H. Aoki, "Science and Medical Applications of Hydroxyapatite". Japanese

Association of Apatite Science. (Tokyo, 1991).

[ 117] L. L. Hench, J. Am. Ceram Soc. 81 ( 1998) 1705.

55

[118] B. D. Ratner, A. S. Hoffman, F. J. Schoen, J. E. Lemons, (editors),

Biomaterials Science (San Diego, California: Academic Press, 1996).

[119] W. Rieger, "High-Tech Ceramics", edited by G. Kostorz (San Diego,

California: Academic Press, 1989) p. 35.

[120] R. J. Young, P.A. Lovell, "Introduction to Polymers", 2nd ed., (Chapman,

Hall, London, 1991) p. 432.

[121] J. Kruger, "Fundamental Aspects of Corrosion of Metallic Implants, in

Corrosion and Degradation of Implant Materials" (eds) B. C. Syrett, A.

Acharya, (ASTM STP 684, p. 107).

[122] L. L. Shreir, R. A. Jarrnen, G. T. Brustein, "Corrosion. Vol. 1:

Metal/Environment Reactions", "Vol. 2: Corrosion Control" 3rd edn

(London: Butterworth Heinemann, 1994).

[123] U. Kamachi Mudali, T. M. Sridhar, B. Raj, Sadhana 28 (2003) 601.

[1?4] M. Sivakurnar, U. Karnachi Mudali, S. Rajeswari, J. Mater. Eng. Perform. 3

(1994) 744.

[125] J. B. Bates, Corrosion 29 (1973) 28.

[126] M. Sivakumar, U. Karnachi Mudali, S. Rajeswari, J. Mater. Sci.: Lett. 13

(1994) 142.

[127] J. Hu, Z. J. Zhano, L. X. Li, Corros. Sci. 35 (1993) 587.

[128] B. C. Syrett, S. S. Wing, Corrosion 11 (1978) 379.

[129] L. L. Hench, E. C. Ethridge, "Biomaterials: An Interfacial Approach",

(Academic Press, New York, 1982).

[130] M. G. Fontana, N. D. Green, "Corrosion Engineering" (New York: Mc

Graw-Hill, 1987).

[131] J. Chiba, H. E. Rubash, K. J. Kim, Clin. Orthop. 300 (1994) 304.

56

[132] D. W. Howie, S. D. Rogers, M.A. McGee, Clin. Orthop. 329S (1996) 217.

[133] T. Visuri, M. Koskenuvouo, Orthopedics 14 (1998) 137.

[134] J.E. Lemons, Clin. Orthop. 235 (1988) 220.

[135] S. D. Cook, K. A. Walsh, R. J. Haddad, Clin. Orthop. Rel. Res. 193 (1985)

271.

[136] G. L. Maistrelli, N. Mahomed, D. G. Fomasier, I. J. Harrington, A.

Binnington, J. Bone joint Sur. 74B (1992) 452.

[137] J. Jansen, E. Ooms, N. Verdonschot, J. Wolke, Orthop. Clin. North. Am.

36 (2005) 89.

[138] D. C. Stanton, J. C. Chou, L. R. Carrasco, J. Oral Maxillofac. Surg. 62

(2004) 235.

[139] W. G. Horstmann, C. C. Verheyen, R. Leemans, Eur. Spine J. 12 (2003)

216.

[140] D. Apelt, F. Theiss, A. 0. El-Warrak, K. Zlinszky, R. B. Wolfisberger, M.

Bohner, Biomaterials 25 (2004) 1439.

[141] M. Nagano, T. Nakamura, T. Kokubo, M. Tanahashi, M. Ogawa,

Biomaterials 17 (1996) 1771.

[142] T. Matsuura, R. Hosokawa, K. Okamoto, T. Kimoto, Y. Akagawa,

Biomaterials 21 (2000) 1121.

[143] L. L. Hench, R. J. Splinter, W. C. Allen, J. Biomed. Mater. Res.

Symposium 2 (1971) 117.

[144] M. Jarcho� J. F. Kay, K. I. Gumaer, R. H. Doremus, H. P. Drobeck, J.

Bioeng. 1 (1977) 79.

57

[145] P. Ducheyne, L. L. Hench, A. Kagan, J. Biomed. Mater. Res. 14 (1980)

225.

[146] H. Oonishi, M. Yamamoto, H. Ishimaru, E. Tsuji, S. Kushitani, M. Aono,

Y. Ukon, J. Bone Joint Surg. (Br) 71 (1989) 213.

[147] Y. T. Sul, C. B. Johansson, Y. Jeong, K. Roser, A. Wennerberg, T.

Albrektsson, J. Mater.Sci.: Mater. Med. 12 (2001) 1025.

[148] E. Fidancevska, G. Ruseska, S. Zafirovski, B. Pavlovski, Science of

Sintering 34 (2002) 241.

[149] B. Feng, J. Weng, J. Zhao, L. He, S. Qi, X. Zhang, Key. Eng. Mater. 240-

242 (2003) 323.

[150] P. Li, Phosphorous Res. Bull. 10 (1999) 232.

[151] X. Chen, Y. Wang, N. Zhao, Y. Zheng, Cailiao Yanjiu Xuebao 17 (2003)

268.

[152] N. Miyata, K. Fuke, Q. Chen, M. Kawashita, T. Kokubo, T. Nakamura,

Biomaterials 25 (2004) 1.

[153] B. G. Cousins, P. J. Doherty, R. L. Willimas, J. Fink, M. J. Garvey, J.

Mater. Sci.: Mater. Med. 15 (2004) 355.

[154] J. Reid, M. Sayer, T. J. Smith, Mater. Res. Soc. Symp. Proc. 2002,

711(Advanced Biomaterials: Characterization, Tissue Engineering and

Comp) 11.

[155] J. Wang, J. De Boer, K. De Groot, J. Dent. Res. 83 (2004) 296.

[156] W. Q. Yan, K. Kawanabe, T. Nakamura, T. Kokubo, Bioceram., Proc. Int.

Symp. Ceram. Med. 10 (1997) 459.

58

[157] L.A. de Sena, M. C. De Andrade, A. M. Rossi, G. de Almeida Soares, J.

Biomed. Mater. Res. 60 (2002) 1.

[158] T. Kasuga, M. Nogami, M. Niinomi, T. Hattori, Biomaterials 24 (2003)

283.

[159] E. Verne, M. Bosetti, C. V. Brovarone, C. Moisescu, F. Lupo, S. Spriano,

M. Cannas, Biomaterials 23 (2002) 3395.

[160] Y. C. Tsui, C. Doyle, T. W. Clyne, Biomaterials 19 (1998) 2015.

[161] M. Inagaki, Y. J. Yokogawa, Key Eng. Mater. 218 (2001) 47.

[162] H. Y. Yasuda, S. Mahara, T. Nishyama, Y. Umakoshi, J. Mater. Sci. Eng. 3

(2002) 29.

[163] J.M. Zhang, C. J. Lin, Z. W. Feng, J. Mater. Sci. Lett. 17 (1998) 1077.

[164] C. K. Wang, J. H. Chem Lin, C. P. Ju, H. C. Og, R. P. H. Chang,

Biomaterials 18 (1997) 1331.

[165] F. Z. Cui, Z. S. Lu, Q. Feng, J. Mater. Sci: Mater. Med. 8 (1997) 403.

[166] Y. Liu, X. Dan, Y. Dan, Y. Du, F. Liu, J. Mater. Sci. 39 (2004) 4031.

[167] T. Kokubo, H. Kushitani, Y. Abe, T. Yamamuro, Bioceramics 2 (1989)

235.

[168] P. Li, C. Ohtuski, T. Kokubo, K. Nakanishi, N. Soga, T. Nakamura, T.

Yamamuro, J. Appl. Biomater. 4 (1993) 221.

[169] F. Barrere, P. Layrolle, C. A. van Blitterswijk, K. de Groot, Mater. Res.

Soc. Symp. Proc. 599 (2000) 135.

[170] P. Habibovik, F. Barrere, C. A. van Blitterswijk, K. de Groot, P. Layrolle,

J. Am. Ceram. Soc. 85 (2002) 517.

59

[171] F. Barrere, C. A. van Blitterswijk, K. de Groot, P. Layrolle, Biomaterials 23

(2002) 1921.

[172] F. Barrere, C. A. van Blitterswijk, K. de Groot, P. Layrolle, Biomaterials 23

(2002) 2211.

[173] C. Y. Yang, R. M. Lin, B. C. Wang, T. M. Lee, E. Chang, Y. S. Hang, P. Q.

Chen, J. Biomed. Mater. Res. 37 (1997) 335.

[174] C. C. Almeida, L.A. Sena, A. M. Rossi, M. Pinto, C. A. Muller, G. A.

Soares, Key Eng. Mater. 254 (2004) 729.

[175] M. Shirkhanzadeh, J. Mater. Sci. Lett. 10 (1991) 1415.

[176] S. Ban, S. Maruno, Biomaterials 19 (1998) 1245.

[177] M. Kumar, H. Dasarathy, C. Riley, J. Biomed. Mater. Res. 45 (1999) 302.

[178] M. C. Kuo, S. K. Yen, Mater. Sci. Eng. C. Biom. Mater. Sens. Syst. 20

(2002) 153.

[179] S. K. Yen, C. M. Lin, Mat. Chem. Phys. 77 (2002) 70.

[180] F. Barrere, C.A. van der Valke, R. A. J. Dalmeijer, G. Meijer, C. A. van

Blitterswijk, K. de Groot, P. Layrolle, J. Biomed. Mater. Res. 66A (2003)

779.

[181] H. Schliephake, D. Scharnweber, M. Dard, S. RoBler, A. Sewing, C.

Hillman, J. Biomed. Mater Res. 64A (2003) 225.

[182] F. Barrere, C. M. van der Valk, G. Meijer, R.A. J. Dalmeijer, K. de Groot,

P. Layrolle, J. Biomed. Mater. Res. 67B (2003) 655.

[183] P. Habibovic, J. P. Li, C. M. van der Valk, G. Meijer, P. Layrolle, C. A.

van Blitterswijk, K. de Groot, Biomaterials 26 (2005) 23.

60

[184] X. X.Wang, W. Yan, S. Hayakawa, K. Tsuru, A. Osaka, Biomaterials 24

(2003) 4631.

[185] X. Lu, Y. Leng, Biomaterials 26 (2005) 1097.

[186] C. Yao, E. B. Slamovich, T. J. Webster, J. I. Qazi, H. J. Rack, Ceramic

Transactions, 159 (Ceramic Nanomaterials and Nanotechnology III, 2005)

239.

[187] Y. Xu, S. LI, Z. Li, Yixue Zazhi 20 (2004) 147.

[188] A. Tampieri, G. Celotti, E. Landi, Anal. Bioanal. Chem. 381 (2005) 568.

[189] L. Yubao, K. de Groot, J. De Wijn, C. P. A. T. Klein, S. V. D. Meer, J.

Mater. Sci. Mat. Med. 5 (1994) 326.

[190] L. He, H. Liu, D. Chen, X. Bai, Surf. Coat. Technol. 160 (2002) 109.

[191] A. Gerlach, B. Vincent, M. Lissac, X. Esnouf, G. Thollet, Biomaterials 14

(1993) 770.

[192] H. Kawamura, A. Ito, T. Muramatsu, S. Miyakawa, N. Ochiai, T. Tateishi,

J. Biomed. Mater. Res. Part A 65 (2003) 468.

[193] M. Ikeuchi, A. Ito, Y. Dohi, H. Ohgushi, H. Shimaoka, K. Yonemasu, T.

Tateishi, J. Biomed. Mater. Res. partA 67 A (2003) 1115.