EFFECT OF ZINC ADDITION ON THE PROPERTIES OF MAGNESIUM
ALLOYS
SAMIR SANI ABDULMALIK
A project report submitted in partial fulfillment of the
requirements for the award of the degree of
Master of Engineering
(Mechanical-Advanced Manufacturing Technology)
Faculty of Mechanical Engineering
University Technology Malaysia
JANUARY 2012
iv
To my mother for her tireless prayers
To Engr Isyaku Jibrin Sani for his Financial Support
v
ACKNOWLEDGEMENT
I would like to say thank you very much to my supervisor Assoc. Prof. Dr.
Mohd Hasbullah Bin Hj. Idris for his wonderful supervision style and encouragement
throughout the project work
My special regards also goes to my mother for her tireless prayers, and to
Engineer Jibring Isyaku Sani for his tremendous financial support
Finally I want to appreciate the effort of all those who have directly or
indirectly contributed to the successful completion of this project work, thank you all.
vi
ABSTRACT
Magnesium alloys are currently used in many structural applications. It is
believed that magnesium and its alloys may also find applications in biomedical
application. In this study, the effects of Zinc (Zn) addition on the properties of
magnesium (Mg) alloys, i.e. Mg–xZn (x = 2, 4, 6, 8, and 10) were investigated. Optical
microscopy, scanning Electron Microscope (SEM), tensile and Vickers hardness testing
were used for the characterization and evaluation of the microstructure and mechanical
properties of the alloys. Electrochemical corrosion measurement was also employed to
determine the corrosion resistance of the alloys. The results show that magnesium alloy
with 6 wt. % zinc content (denoted as Mg- 6Zn) shows good corrosion resistance and
mechanical properties).
vii
ABSTRAK
Pada masa ini aloi magnesium (Mg) telah digunakan dalam pelbagai aplikasi
struktur. Dipercayai bahawa magnesium dan aloinya telah digunakan dalam bidang
bioperubatan. Dalam kajian ini, kasan pertambahan zink (Zn) (2, 4, 6, 8 dan 10% berat)
tehadap sifat mekanikal dan kakisan aloi magnesium, Mg-xZn telah dikaji. Analisis
menggunakan mikroscop optik, Scanning Electron Mikroscopy (SEM),ujian ketegangan
dan kekerasan Vickers telah digunakan bagi pencirian dan penilaian mikrostructur dan
sifat mekanikal aloi yang dikaji. Ujian kakisan electrokimia juga telah digunakan untuk
menilai sifat rintangan kakisan aloi. Keputusan ujikaji menunjukan bahawa aloi
magnesium dengan kandungan 6% berat zink (diwakili dengan Mg-6Zn) memberikan
sifat kakaisan dan mekanikal yang baik.
viii
TABLE OF CONTENTS
CHAPTER
1
2
TITLE
PAGE
DECLARATION
ii
DEDICATION
iv
ACKNOWLEDGMENT
v
ABSTRACT
vi
ABSTRAK
vii
LIST OF CONTENTS
viii
LIST OF TABLES
xii
LIST OF FIGURES
xiii
LIST OF APPENDICES
xv
INTRODUCTION
1.1
Background
1
1.2
Statement of Problem
3
1.3
Objectives
3
1.4
Scopes
4
LITRETURE REVIEW
2.1
Overview of Biomaterials
5
2.1.1 Uses for Biomaterials
6
ix
2.1.1.1 Orthopedics
6
2.1.1.2 Cardiovascular Applications
7
2.1.1.3 Ophthalmic
7
2.1.1.4 Dental Applications
7
2.1.2 Types of Biomaterials
2.1.2.1 Metallic materials
8
2.1.2.2 Polymers
9
2.1.2.3 Ceramics
9
2.1.2.4 Composites
10
2.1.3 Natural Biomaterials
11
2.1.4 Application of Biomaterials
11
2.2 Natural Bone
2.2.1 Desirable Properties of Artificial Bone Material
12
13
2.2.1.1 Body Condition
13
2.2.1.2 Mechanical Properties
15
2.2.1.3 Corrosion Resistance
15
2.3 Conventional Metallic Materials Used For Medical Devices
2.4
7
16
2.3.1 Stainless steels
17
2.3.2 Cobalt-Base Alloys
17
2.3.3 Titanium and Titanium-Base Alloys
18
Magnesium
18
2.4.1 Properties of Pure Magnesium
19
2.4.2
Melting and casting of magnesium
20
2.4.2.1 Melting
20
x
2.4.3
2.5
2.6
3
2.4.2.2 Casting and working of magnesium
22
Magnesium Alloys
22
2.4.3.1 Common Alloying Elements
23
2.4.3.1.1 Aluminum
23
2.4.3.1.2 Calcium
23
2.4.3.1.3 Manganese
23
2.4.3.1.4 Rare Earths
24
2.4.3.1.5 Zinc
24
Zinc Metal
24
2.5.1 Zinc Biological role
24
Researched Biodegradable Magnesium Alloys
25
RESEARCH METHODOLOGY
3.1
Introduction
28
3.2
Research Design
30
3.2.1
30
Casting
3.2.2. Microstructural Characterization
33
3.2.3
Hardness Test
35
3.2.4
Tensile Test
37
3.2.5 Electrochemical Measurement
4
38
RESULTS AND DISCUSSION
4.1
Selection of optimum zinc addition
39
4.1.1 Nominal Composition Analysis
39
xi
4.1.2
Microstructural Characterization
4.1.3 Hardness Test
43
4.1.4
44
Tensile Test
4.1.5 Corrosion Electrochemical Test
5
40
CONCLUSION
46
47
REFERENCES
48
APENDIX A-C
51
xii
LIST OF TABLES
TABLE NO.
TITLE
PAGE
2.1
Example of Medical and Dental Material and their Applications
2.2
Example of Polymers used as Biomaterials
2.3
Example of Biomaterial Ceramics
2.4
Summary of the mechanical properties and porosity of human
bone
2.5
Raw Materials for Magnesium Production
4.1
Nominal chemical composition of the Mg-Zn alloys
4.2
The Tensile strength, Yield, and Elastic Modulus value
for Mg-Zn alloys
8
9
10
15
19
39
45
xiii
LIST OF FIGURES
FIGURE NO.
TITLE
PAGE
2.1
Implant material requirements in orthopedic applications
6
2.2
Hip joint replacement
12
2.3
Details of the bone structure
13
2.4
Closed packed structure of pure magnesium
20
3.1
Flowchart showing the summary of research methodology
29
3.2
(a) Magnesium Ingot (b) Pure Zinc
30
3.3
(a) Mg-Zn Melting, (b) Pouring into steel mold,
(c) Designed Mold, (e) cast sample (f) the mold used
3.4
(a) Olympus BX60, (b) Philips XL 40, (c) Supra 35VP,
used for the characterization of the microstructure
3.5
37
Electrochemical test (Parstat-2263) set up used for the corrosion
Measurement
4.1
36
Instron universal tensile testing machine used in the tensile
testing of the samples
3.7
34
(a) Matsuzawa DVK-2 used for the hardness testing
(b) location of the test on the sample
3.6
31
38
Microstructure of the as cast (a) pure magnesium, (b) Mg-2Zn,
(c) Mg-4Zn, (d) Mg-6Zn, (e) Mg-8Zn, (f) Mg-10Zn
40
xiv
4.2
FE-SEM micrographs of (a) Mg-8Zn alloy,
(b) Mg-10Zn Alloy, (c) (Mg, Zn)-containing phase in the grain,
(d) (mg, Zn)-containing phase at the grain boundary
4.3
EDS analysis of the secondary phases (a) on the grain,
(b) at the grain boundary
4.4
43
The Tensile strength value of Mg-Zn alloys as a function
of zinc addition.
4.6
42
The hardness value of Mg-Zn alloys as a function
of zinc addition
4.5
41
45
Electrochemical polarization curves of Mg-Zn alloys
under investigation
46
xv
LIST OF APPENDICES
APPENDIX
TITLE
PAGE
A
Compositional Analysis of the as- cast Samples
51
B
The Stress/Strain Graphs for the alloys Samples
61
C
Polarization Curves of the Samples
66
1
CHAPTER 1
INTRODUCTION
1.1
Background
Biomaterial implants are used as a replacement of a bone part or as a support in
the healing process. Replacement of a bone part requires implants to stay in the body
permanently, while support only requires that the implant remain in the body for a
shorter period. When permanent implant is used for a temporary application, additional
surgeries are required to remove these devices after the healing process. Thus, removal
process increases the patient grim and cost of health care. In contrast, biodegradable
materials require no additional surgeries for removal as they dissolve after the healing
process is complete. This also eliminates the complications associated with the longterm presence of implants in the body. Finally, after these materials degrade within the
body, it is important that the body can metabolized the degradation products, and thus
are bioabsorbable.
The first materials to be used as commercial biodegradable and bioabsorbable
implant materials were polymers. The most commonly and earliest used absorbable
materials include polyglycolic acid (PGA), poly-lactic acid (PLA), and poly-dioxanone
(PDS). However, low mechanical properties and radiolucency are the limitation with
these materials. Applications of polymeric materials in load-bearing and tissue
2
supporting applications is severely restricts due to low strength, because the mechanical
needs of the body required a greater amount of material.
Metals due to their relatively high strength and fracture toughness possesses
desirable mechanical properties, however, most of the metals are biologically toxic.
Studies revealed that conventional implant, like cobalt, stainless, chromium, and nickelbased alloys produce corrosion products, which are harmful to the human body [1] [2]
[3] [4].
Magnesium and its alloys are biodegradable metals and exhibit improved
mechanical properties and corrosion resistance. However most of the reported
biomedical magnesium alloys contain aluminum and/or rare earth (RE) elements. It is
well known that Al and rare earth elements are harmful to neurons, osteoblasts, and also
associated with dementia and could lead to hepatotoxicity. Consequently, Al and RE are
unsuitable alloying elements for biomedical magnesium materials, particularly when
they are above certain levels [5]
Pure magnesium was indicated as suitable candidate for temporary implant;
however, the major drawback of Mg is its low corrosion resistance which results to low
mechanical strength in the physiological environment. Alloying elements can be added
to increase the strength of pure Mg but alloying elements should be selected carefully to
maintain the Mg’s biocompatibility.
With the purpose of searching for suitable alloying elements for biomedical
magnesium alloys, researchers demonstrated that Calcium (Ca), Manganese (Mn), and
Zinc (Zn) could be appropriate candidates. Zinc is one of the essential elements in
human body that also provide mechanical strengthening in Mg-based alloys.
Zinc can improve the corrosion resistance and mechanical properties of
magnesium alloys, Zinc additions increase the strength of Mg-based alloys primarily
through precipitation strengthening. Furthermore, zinc is one of the most abundant
3
nutritionally essential elements in the human body, and has basic safety for biomedical
applications [6] [7].
1.2
Statement of Problem
The mechanical properties and corrosion resistance of magnesium alloys must
be sufficiently investigated for medical application. Magnesium is essential to human
metabolic functions and is the fourth most abundant cation in human body. In vitro
cytotoxicity of pure magnesium metal showed positive cell proliferation and viability
with no sign of growth inhibition. The fracture toughness of magnesium is greater than
that of ceramics, but pure magnesium corrodes too quickly in the physiological
environment (pH 7.4–7.6), losing mechanical integrity before tissue healing. In an effort
to maintain the mechanical integrity, and biocompatibility, more alloying compositions
are necessary.
1.3
Objectives
The objectives of this project are:
1. To establish optimum material composition Mg-Zn
2. To establish the effect of Zinc addition on the properties of Mg alloy as
biodegradable material
4
1.4
Scopes
This project was conducted within the following boundaries:
1. Mg-Zn alloys was prepared and cast using gravity die casting process
2. The effect of zinc addition was characterized and measured through:
(a) Microstructure observation
(b) Mechanical properties test, and
(c) Electrochemical corrosion tests
5
CHAPTER 2
LITRATURE REVIEW
2.1
Overview of Biomaterials
A biomaterial is any synthetic material that is used to replace or restore function
to a body tissue and is continuously or intermittently in contact with body fluids.
Exposure to body fluids usually implies that the biomaterial is placed within the interior
of the body, and this places several strict restrictions on materials that can be used as a
biomaterial [8].
Biomaterial must be biocompatible; it should not elicit an adverse response
from the body, and vice versa. Additionally, it should be nontoxic and noncarcinogenic.
These requirements eliminate many engineering materials that are available. Next, the
biomaterial should possess adequate physical and mechanical properties to serve as
augmentation or replacement of body tissues. For practical use, a biomaterial should be
amenable to being formed or machined into different shapes, have relatively low cost,
and be readily available. Various material requirements that must be met for successful
total joint replacement are listed in Figure 2.1 [9].
6
Figure 2.1: Implant material requirements in orthopedic applications. [9]
2.1.1
Uses for Biomaterials
Biomaterials are primarily used to replace hard or soft tissues that have become
destroyed or damaged through some pathological process. As a result of these
circumstances, it may be possible to remove the damaged tissue and replace it with some
suitable synthetic material [10]. Listed below are some common uses of biomaterials.
2.1.1.1 Orthopedics
Orthopedic implant devices are one of the most prominent application areas for
biomaterials. It has been possible to replace
joints, such as the hip, knee, shoulder,
ankle, and elbow, and the pains resulted can be considerable, since the introduction of
anesthesia, antisepsis, and antibiotics. The relief of pain and restoration of mobility is
well known to hundreds of thousands of patients.
7
2.1.1.2 Cardiovascular Applications
Problems arose with heart valves and arteries can be successfully treated with
implants. The heart valves sometimes fails to either fully opening or fully closing,
meaning to say the valve is affected with disease, the diseased valve can be replaced
with a variety of substitutes.
2.1.1.3 Ophthalmic
The tissues of the eye suffer from several diseases, leading to reduced vision and
eventually, blindness. Cataracts, for example, cause cloudiness of the lens. This may be
replaced with a synthetic (polymer). As with intraocular lenses, biomaterials are used to
preserve and restore vision
2.1.1.4 Dental Applications
Within the mouth, both the tooth and supporting gum tissues can be readily
destroyed by bacterially controlled diseases. Teeth in their entirety and segments of teeth
both can be replaced and restored by a variety of materials.
2.1.2
Types of Biomaterials
In general, synthetic biomaterials used for implants can be categorized as:
metals, polymers, ceramics, and composites [8]
8
2.1.2.1 Metallic materials
Metallic materials are the most widely used for load-bearing implants. For
instance, some of the most common orthopedic surgeries involve the implantation of
metallic implants. These range from simple wires and screws to fracture fixation plates
and total joint prostheses (artificial joints) for hips, knees, shoulders, ankles, and so on.
In addition to orthopedics, metallic implants are used in maxillofacial surgery,
cardiovascular surgery, and as dental materials. Although many metals and alloys are
used for medical device applications, the most commonly employed are stainless steels,
commercially pure titanium and titanium alloys, and cobalt-base alloys (Table 2.1).
Table 2.1: Example of Medical and Dental Material and their Applications
9
2.1.2.2 Polymers
Polymers are used in medicine as biomaterials. Their applications range from
facial prostheses to tracheal tubes, from kidney and liver parts to heart components, and
from dentures to hip and knee joints. Polymeric materials are also used for medical
adhesives and sealants and for coatings that serve a variety of functions. Example of
polymer material is shown in Table 2.2.
Table 2.2: Example of Polymers used as Biomaterials
2.1.2.3 Ceramics
Traditionally, ceramics have seen widescale use as restorative materials in
dentistry. These include materials for crowns, cements, and dentures. However, their use
in other fields of biomedicine has not been as extensive, compared to metals and
polymers. For example, the poor fracture toughness of ceramics severely limits their use
for load-bearing applications. Some ceramic materials are used for joint replacement and
bone repair and augmentation as shown in the Table 2.3.
10
Table 2.3: Example of Biomaterial Ceramics
Ceramics and glasses
Applications
Alumina
Join replacement, dental implants
Zirconia
Join replacement
Calcium phosphate
Bone repair and augmentation, surface
coatings on metals
Bioactive glasses
Bone replacement
Porcelain
Dental restoration
Carbons
Heart valves, percutaneous devices, dental
implants
2.1.2.4 Composites
Composites biomaterials are used in the field of dentistry as restorative
materials or dental cements. Although carbon-carbon and carbonreinforced polymer
composites are of great interest for bone repair and joint replacement because of their
low elastic modulus levels, these materials have not displayed a combination of
mechanical and biological properties appropriate to these applications. Composite
materials are, however, used extensively for prosthetic limbs, where their combination
of low density/weight and high strength make them ideal materials for such applications.
11
2.1.3
Natural Biomaterials
One of the advantages of using natural materials for implants is that they are
similar to materials familiar to the body. Natural materials do not usually offer the
problems of toxicity often faced by synthetic materials. Also, they may carry specific
protein binding sites and other biochemical signals that may assist in tissue healing or
integration. The problem with the natural materials is that they can be subjected to
immunogenicity, and their tendency to denature or decompose at temperatures below
their melting points. This severely limits their fabrication into implants of different sizes
and shapes. An example of a natural material is collagen, which exists mostly in fibril
form, has a characteristic triple-helix structure, and is the most prevalent protein in the
animal world [8].
2.1.4
Application of Biomaterials
Total joint replacement is widely regarded as the major achievement in
orthopedic surgery in the 20th century. Arthroplasty, or the creation of a new joint, is the
name given to the surgical treatment of degenerate joints aimed at the relief of pain and
the restoration of movement. This has been achieved by excision, interposition, and
replacement arthroplasty and by techniques that have been developed over
approximately 180 years. Hip arthroplasty generally requires that the upper femur (thigh
bone) be replaced and the mating pelvis (hip bone) area be replaced or resurfaced. As
shown in Figure 3, a typical hip prosthesis consists of the femoral stem, a femoral ball,
and a polymeric (ultrahigh molecular weight polyethylene, or (UHMWPE)) socket (cup)
with or without a metallic backing [9].
12
Figure 2.2: Hip joint replacement
2.2
Natural Bone
Bone is a composite type substance containing calcium, phosphate, magnesium
and collagen. Structurally bone is divided into five parts, namely:
1. Periostium
2. Compact bone
3. Spongy bone
4. Bone marrow, and
5. Epiphyseal plate
Bones are rigid and elastic in nature. Major percent of the bone is hydroxyapatite and
another small percent of carbonate is present in human bone.
13
Figure 2.3: Details of the bone structure [11]
Mineral substances in the bone like calcium, phosphate and magnesium make the bone
as rigid substance and collagen makes it as elastic substance. During bone development
stage mineral substance are converted into apatite minerals from crystallographicall
amorphous [11] [12].
2.2.1
Desirable Properties of Artificial Bone Material
2.2.1.1 Body Condition
Temperature conditions are not extreme for the materials used, body
temperatures being a little less than 38°C (98.6°F). However, the chemical physiological
environment and biomechanical environment can be extreme. For structural implants
used to repair the hip, it is estimated that the average nonactive person may place 1 to
14
2.5 × 106 cycles of stress on his or her hip in a year. For a person 20 to 30 years of age,
with a life expectancy of 70 to 80 years, that is the equivalent of approximately 108
cycles of loading in a lifetime. The actual loads and cycles are a function of the weight
and activity level of the person, but the need for longtime cyclic capability in fatigue is
obvious. Other applications in the body also impart many millions of fatigue cycles to
the device or component implanted.
In considering the parameters of materials for intracorporeal applications,
several factors are of major importance. It is generally agreed that the material must:
i.
Be nontoxic and noncarcinogenic, cause little or no foreign-body reaction, and be
chemically stable and corrosion resistant. This is known as biocompatibility.
ii.
Be able to endure large and variable stresses in the highly corrosive environment
of the human body
iii.
Be able to be fabricated into intricate shapes and sizes
Many structural applications of materials in the body require that the
replacement material fit into a space perhaps only one-fourth the area of the part being
permanently or temporarily replaced or assisted. Consequently, the implant may have to
withstand loads up to 16 or more times that which the human bone must withstand. In
restorative dentistry, high compressive biting forces are combined with large
temperature changes and acidity to produce a challenging environment. It is clear that
there can be very great mechanical loading demands on biomaterials used for structural
purposes [13]
15
2.2.1.2 Mechanical Properties
Mechanical properties of artificial bone material should be similar to the natural
bone. Rejection of artificial implants due to mismatch in mechanical property between
natural bone and implant is known as biomechanical incompatibility.
Important mechanical properties are tensile strength, hardness and modulus of
elasticity. Artificial bone material considered for implant should have high strength and
low modulus of elasticity to match the property of natural bone [14].
Table 2.4: Summary of the mechanical properties and porosity of human bone [14]
Bone
Cortical
Bone
Cancellous
bone
Compressive
Strength
(MPa)
130-180
Flexural
Strength
(MPa)
135-193
Tensile
Strength
(MPa)
50-150
Modulus
(MPa)
Porosity
(%)
12-18
5-13
4-12
NA (Not
1-5
0.1-0.5
30-90
available
2.2.1.3 Corrosion Resistance
Corrosion has been a major determining factor in the selection of materials for
use in the body environment. The first requirement for any material whether a
metal/alloy, ceramic, or polymer to be placed in the body is that it should be
biocompatible and not cause any adverse reaction in the body. The material must
withstand the body environment and not degrade to the point that it cannot function in
the body as intended. For example, metals used in the cardiovascular system must be
nonthrombogenic, and, in general, the more electronegative the metal with respect to
blood, the less thrombogenic the metal will be. For a material to be considered resistant
16
to corrosion in the body, its general corrosion rate usually must be less than, 0.01 mil/yr
(0.00025 mm/yr).
In vitro electrochemical measurements can be conducted in controlled
environments, and these techniques provide methods of determining the basic corrosion
reactions necessary for predicting the corrosion behavior of materials and for screening
and characterizing materials intended for use in surgical applications [15].
2.3
Conventional Metallic Materials Used For Medical Devices
Metals have been successfully used as biomaterials for many years. Besides
orthopedics, there are other markets for metallic implants and devices, including oral
and maxillofacial surgery (e.g., dental implants, craniofacial plates and screws) and
cardiovascular surgery (e.g., parts of pacemakers, defibrillators, and artificial hearts;
balloon catheters; valve replacements; stents; and aneurysm clips). Surgical instruments,
dental instruments, needles, staples, and implantable drug pump housings are also made
from metallic materials.
For structural applications in the body (e.g., implants for hip, knee, ankle,
shoulder, wrist, finger, or toe joints), the principal metals are stainless steels, cobalt-base
alloys, and titanium-base alloys. These metals are popular primarily because of their
ability to bear significant loads, withstand fatigue loading, and undergo plastic
deformation prior to failure. Other metals and alloys employed in implantable devices
include commercially pure titanium (CP-Ti), shape memory alloys (alloys based on the
nickel-titanium binary system), zirconium alloys, tantalum (and, to a lesser extent,
niobium), and precious metals and alloys [16].
17
2.3.1
Stainless steels
Stainless steels are iron-base alloys that contain a minimum of 10.5% Cr, the
amount needed to prevent the formation of rust in unpolluted atmospheres. Stainless
steels used for implants are suitable for close and prolonged contact with human tissue
(i.e., warm, saline conditions). Specific requirements for resistance to pitting and crevice
corrosion and the quantity and size of nonmetallic inclusions apply to implant-grade
stainless steels.
Austenitic stainless steels are popular for implant applications because they are
relatively inexpensive, they can be formed with common techniques, and their
mechanical properties can be controlled over a wide range for optimal strength and
ductility. Stainless steels for implants are wrought alloys (i.e., they are fabricated by
forging and machining). Austenitic stainless steels are not sufficiently corrosion resistant
for long-term use as an implant material. They find use as bone screws, bone plates,
intramedullary nails and rods, and other temporary fixation devices.
Recently, other stainless steels ‘nitrogen-strengthening stainless steel’ has been
developed and standardized that have increased corrosion resistance and improved
mechanical properties. Nitrogen-strengthened alloys are being used for bone plates, bone
screws, spinal fixation components, and other medical components. Nitrogenstrengthened stainless steels have better crevice and pitting corrosion resistance.
2.3.2
Cobalt-Base Alloys
Cobalt-base alloys were first used in the 1930s. The Co-Cr-Mo alloy Vitallium
was used as a cast dental alloy and then adopted to orthopedic applications starting in the
1940s. The corrosion of cobalt-chromium alloys is more than an order of magnitude
greater than that of stainless steels, and they possess high mechanical property
18
capability. Although cobalt alloys were first used as cast components, wrought alloys
later came into use.
2.3.3
Titanium and Titanium-Base Alloys
Titanium and its alloys used for implant devices have been designed to have
excellent biocompatibility, with little or no reaction with tissue surrounding the implant.
Titanium derives it corrosion resistance from the stable oxide film that forms on its
surface, which can reform at body temperatures and in physiological fluids if damaged.
Increased use of titanium alloys as biomaterials is occurring due to their lower modulus,
superior biocompatibility, and enhanced corrosion resistance when compared to more
conventional stainless steels and cobaltbase alloys.
2.4
Magnesium
Magnesium always appears in nature in ionic form with the following electron
arrangement: 1S22S22P63S2.This arrangement is characterized by the low ionization
energies relative to the two most external electrons, which are at the 3S level. The low
standard reduction potential of magnesium is the reason why no metallic magnesium is
found in nature:
Mg2+ + 2e– = Mg E0 = –2.375 V
The raw materials for the production of magnesium come from different magnesium
sources. In all cases they will be accompanied by additional materials, depending on
their source.
19
Table 2.5: Raw Materials for Magnesium Production
Material
Chemical formula
Magnsite
MgCO3
Dolomite
MgCO3·CaCO3
Bischofite
MgCL2·6H2O
Carnallite
MgCL2·KCL·6H2O
Serpentine
3MgO·2SiO2·2H2O
Sea water
Mg2+(aq)
2.4.1
Properties of Pure Magnesium
Magnesium is classified as an alkaline earth metal. It is found in Group 3 of the
periodic table, and has the atomic properties as:
i.
Element Symbol
Mg
ii.
Atomic Number
12
iii.
Atomic Weight
24.3050
iv.
Atomic Diameter 0.320 nm
v.
Atomic Volume
14.0 cm3/mol
Lattice parameters of pure magnesium estimated at room temperature are: a = 0.32092
nm and c = 0.52105 nm. The c/a ratio is 1.6236 which is close to the ideal value of
1.633. Therefore, magnesium is considered as perfectly closed packed [17].
20
Figure 2.4: Closed packed structure of pure magnesium
The density of magnesium at 20°C is 1.738 g/cm3.At the melting point of 650°C
reduced to 1.65 g/cm3, on melting there is an expansion in volume of 4.2%. At higher
temperature volume diffusion is very important, especially at T > 0.6 T m, where Tm is
the absolute melting point. At lower temperatures pipe diffusion, i.e., diffusion along
dislocation cores, may become more significant. Grain boundary diffusion plays a role
in polycrystals because the grain boundary acts as a low energy channel for the
movement of atoms.
The thermal conductivity of pure magnesium measured at elevated temperatures
decreases with increasing temperature. At very low temperatures the thermal
conductivity exhibits high values [17]
2.4.2
Melting and casting of magnesium
2.4.2.1 Melting
It is usual for magnesium to be melted in mild steel crucibles for both the
alloying and refining or cleaning stages before producing cast or wrought components.
21
Unlike aluminium and its alloys, the presence of an oxide film on molten magnesium
does not protect the metal from further oxidation. On the contrary, it accelerates this
process. Melting is complete at or below 650 °C and the rate of oxidation of the molten
metal surface increases rapidly with rise in temperature such that, above 850 °C, a
freshly exposed surface spontaneously bursts into flame. Consequently, suitable fluxes
or inert atmospheres must be used when handling molten magnesium and its alloys.
For many years, thinly fluid salt fluxes were used to protect molten magnesium
which were mixtures of chlorides such as MgCl2 with KCl or NaCl. However, the
presence of the chlorides often led to problems with corrosion when the cast alloys were
used in service [19].
Cover gases (e.g. SO2 or argon), or a mixture of an active gas diluted with CO2,
N2, replaced the used of the salt, also sulphur hexafluoride (SF6) was widely accepted as
the active gas because it is non-toxic, odourless, colourless, and effective at low
concentrations. But the disadvantage of sulphur hexafluoride SF6 is, however, relatively
expensive, and is now realised to be a particularly potent greenhouse gas with a socalled Global Warming Potential (GWP) of 22, 000 to 23, 000 on a 100 year time
horizon.
As a result of that efforts are therefore being made to find other active gases
containing fluorine and one promising alternative is the organic compound HFC 134a
(1,1,1,2-tetrafluoroethane) that is readily available worldwide because of its use as a
refrigerant gas. It is also less expensive than SF6. HFC 134a has a GWP of only 1600,
and an estimated atmospheric lifetime of 13.6 years compared with 3, 200 years for SF6.
Moreover less is consumed on a daily basis so that the overall potential to reduce
greenhouse gas emissions is predicted to be 97% [19].
22
2.4.2.2
Casting and working of magnesium
Most magnesium components are now produced by high-pressure die casting
machine. Cold chamber machines are used for the largest castings and molten shot
weights of 10 kg or more can now be injected in less than 100 ms at pressures that may
be as high as 1500 bar. Hot chamber machines are used for most applications and are
more competitive for smaller sizes due to the shorter cycle times that are obtainable.
A reported disadvantage with high pressure die castings is that they may contain
relatively high levels of porosity. This restricts opportunities for using heat treatment to
improve their properties because exposure to high temperatures may cause the pores to
swell and form surface blisters.
Sand castings and low pressure permanent mould castings generally contain
less porosity and are used to produce components having more complicated shapes.
They can then be heat treated if the alloys respond to age hardening. With permanent
mould casting, turbulence can be minimized by introducing the molten metal into the
bottom of the mould cavity, under a controlled pressure, thereby allowing unidirectional
filling of the mould [19].
2.4.3
Magnesium Alloys
Magnesium is readily available commercially with purities exceeding 99.8%
but it is rarely used for engineering applications without being alloyed with other metals.
The fact that its atomic diameter (0.320 nm) is such that it enjoys favourable size factors
with a diverse range of solute elements.
23
2.4.3.1 Common Alloying Elements
2.4.3.1.1 Aluminum
Aluminum is the most commonly used alloying element, and the maximum
solubility is 11.5 at % (12.7 mass %) and alloys in excess of 6 mass % can be heat
treated. Aluminum improves strength, the optimum combination of strength and
ductility being observed at about 6%.Alloys is readily castable.
2.4.3.1.2 Calcium
Alloying with calcium is becoming more common in the development of
cheap creep resistant alloys. Ca can act as deoxidant in the melt or in subsequent heat
treatment. It improves the roll ability of sheet but >0.3 mass % can reduce the weld
ability.
2.4.3.1.3 Manganese
Manganese is usually not employed alone but with other elements, e.g., Al. It
reduces the solubility of iron and produces relatively innocuous compounds. It increases
the yield strength and improves salt water corrosion resistance of Mg-Al and Mg-Al-Zn
alloys. Binary alloys (M1A) are used in forgings or extruded bars. The maximum
amount of manganese is 1.5–2 mass %.
24
2.4.3.1.4 Rare Earths
Rare Earths are added to magnesium alloys to improve the high temperature
strength, and creep resistance; they are usually added as Mischmetal.
2.4.3.1.5 Zinc
Zinc is one of the commonest alloying additions. It is used in conjunction with
Al (e.g., AZ91 or with zirconium, thorium or rare earths) [18].
2.5
Zinc Metal
Zinc is, in some respects, chemically similar to magnesium, because its ion is of
similar size and its only common oxidation state is +2. Zinc is the 24th most abundant
element in the Earth's crust and has five stable isotopes [19].
2.5.1
Zinc Biological role
Zinc is an essential trace element, necessary for plants, animals, and
microorganisms. Zinc is found in nearly 100 specific enzymes (other sources say 300),
serves as structural ions in transcription factors and is stored and transferred in
metallothioneins. It is "typically the second most abundant transition metal in
organisms" after iron and it is the only metal which appears in all enzyme classes.
In proteins, Zn ions are often coordinated to the amino acid side chains of
aspartic acid, glutamic acid, cysteine and histidine. The theoretical and computational
25
description of this zinc binding in proteins (as well as that of other transition metals) is
difficult. There are 2–4 grams of zinc distributed throughout the human body. Most zinc
is in the brain, muscle, bones, kidney, and liver, with the highest concentrations in the
prostate and parts of the eye. Semen is particularly rich in zinc, which is a key factor in
prostate gland function and reproductive organ growth.
In humans, zinc plays "ubiquitous biological roles". It interacts with "a wide
range of organic ligands", and has roles in the metabolism of RNA and DNA, signal
transduction, and gene expression. It also regulates apoptosis. A 2006 study estimated
that about 10% of human proteins potentially bind zinc, in addition to hundreds which
transport and traffic zinc; a similar in silico study in the plant Arabidopsis thaliana found
2367 zinc-related proteins [21].
In the brain, zinc is stored in specific synaptic vesicles by glutamatergic neurons
and can "modulate brain excitability". It plays a key role in synaptic plasticity and so in
learning. However it has been called "the brain's dark horse" since it also can be a
neurotoxin, suggesting zinc homeostasis plays a critical role in normal functioning of the
brain and central nervous system [21].
2.6
Researched Biodegradable Magnesium Alloys
For the purpose of searching for suitable alloying elements for biomedical
magnesium alloys, researchers exploited the in vivo and the in vitro behavior of
magnesium alloys.
Yizao Wan [22] in his research work named preparation and characterization of
a new biomedical magnesium–calcium alloy. Demonstrated that 0.6 wt % calcium
content improved corrosion and mechanical properties of magnesium and the alloy Mg0.6Ca shows good potential as a new biomedical material.
Hui Du [23] researched on the effects of the addition of Zn element on the
properties of Mg–3Ca. He pointed out that the element Zn could improve both tensile
26
strength and elongation of Mg–3Ca alloys, and attributed that the presence of
Ca2Mg6Zn3 phase found in the alloy mainly contributes to these improvement.
Jun Wang [24] investigates the Microstructure and corrosion properties of as
sub-rapid solidification of Mg–Zn–Y–Nd alloy in dynamic simulated body fluid for
vascular stent application. The research shows that as sub-rapid solidification can
improve the corrosion resistance of Mg–Zn–Y–Nd alloy in dynamic SBF due to grain
refinement.
Yuncang Li [25] exploded the effects of calcium (Ca) and yttrium (Y) on the
microstructure, mechanical properties, corrosion behavior and biocompatibility of
magnesium (Mg) alloys. Results of the investigation indicated that Ca content can
enhances the compressive strength, elastic modulus and hardness of the Mg–Ca alloys,
but deteriorates the ductility, corrosion resistance and biocompatibility of the Mg–Ca
alloys. Also revealed that yttrium addition increases ductility; but decreases the
compressive strength, hardness, corrosion resistance and biocompatibility of the alloy
Mg–1Ca–1Y.
Erlin Zhang [26] research on the in vivo degradation of magnesium alloy
implant highlighted that rapid degradation of magnesium implant was observed in the
marrow channel than in the cortical bone. Also shown that the degradation of the
magnesium implant in the blood caused little change to blood composition but no
disorder to liver or kidneys
E. Aghion G. Levy [27] evaluated the effect of 0.4% Ca on the in vitro
corrosion behavior of Mg–1.2% Nd–0.5% Y–0.5% Zr in a simulated physiological
environment. He outlined that 0.4% Ca has a beneficial effect on the corrosion resistance
of the tested alloy, and attributed this to the effect of calcium, which reduces oxidation
in the molten condition and consequently improves the soundness of the obtained
casting, E. Aghion result also shown that the addition of calcium has a damaging effect
on the stress corrosion behavior in terms of reduced ultimate tensile strength and
27
ductility of the alloy, and this was mainly due to the embrittlement effect of calcium that
was generated by the formation and distribution of Mg2Ca phase at grain boundaries.
Zijian Li [28] in his work, research on the development of binary Mg-Ca alloys
for use as biodegradable materials reveals that controlled calcium content and processing
treatment can lead to the improvement in tensile strength and corrosion properties of the
alloy. Also highlighted that Mg-1Ca alloy show an acceptable biocompatibility as a new
kind of biodegradable implant material.
Yingwei Song [29] explores the in vitro corrosion behaviors of the
biodegradable AZ31 in simulated body fluid (SBF), pointing out that some protective
film layer was formed on the surface of AZ31 and had perfect biocompatibility.
X.N. GU [30] investigates the Corrosion fatigue behaviors of two biomedical
Mg alloys AZ91D and WE43 – In simulated body fluid. Demonstrated that die-cast
AZ91D alloy indicated a lower fatigue limit than that observed for extruded WE43 alloy
.
Yin Dongsong [31] demonstrated that Zinc content (3% wt) refined the
microstructure, and improved the mechanical properties of Mg-Mn alloy.
Shaoxiang Zhang [32] explored in vitro and in vivo potentials of Mg-Zn alloy
and point out the alloy shown suitable mechanical properties for implant application and
also good corrosion resistance.
28
CHAPTER 3
RESEARCH METHODOLOGY
3.1
Introduction
The experimental work in this project was to prepare the magnesium zinc alloys
and to study the effect of zinc addition on the microstructure, mechanical and corrosion
properties of Mg-Zn alloys. Figure 3.1.shown the general flowchart of the experimental
procedures.
29
Start
Prepare
Mg-Zn
Prepare mold
Gravity die mold
Melting at 7000C7500C in mild
steel crucible
Cast at 7300C
into steel mold
Microstructure
characterization
Mechanical Test
Tensile/Hardness
test
Result/Discussion
Stop
Figure 3.1: Flowchart showing the summary of research methodology
30
3.2
Research Design
3.2.1
Casting
Pure magnesium ingot (99.99 wt %) Figure 3.2(a) and, pure zinc ingot (99.995
wt. %), Figure 3.2(b), were used in this experiment. Melting process was carried out in a
high frequency induction furnace (inductothem) Figure 3.3a using a mild steel crucible
under the argon gas atmosphere. After magnesium metal was melted at about 6500C,
pure zinc ingot was added. After all these materials were melted completely and
superheat to around 7500C, the melt was then cast into a steel mold at 7300C Figure
3.3b. Figure 3.3(c, d and e) shows the schematic mold design, as cast sample, and the
original mold used in the casting process.
(a)
(b)
Figure 3.2: (a) Magnesium Ingot (b) Pure Zinc
31
(a)
(b)
32
(c)
(d)
33
Figure 3.3: (a) Mg-Zn Melting, (b) Pouring into steel mold, (c) Designed Mold, (e) Cast
Sample, (f) the mold used
3.2.2.
Microstructural Characterization
Characterization of the microstructure and phases of the Mg-Zn alloys was
conducted using optical microscope (Olympus BX60) and scanning electron microscope
(Philips XL 40) equipped with energy dispersive spectroscopy (EDS). And carls zeiss
supra 35VP (FE-SEM), Shown in figure 3.4.
34
(a)
(b)
35
(c)
Figure 3.4: (a) Olympus BX60, (b) Philips XL 40, (c) Supra 35VP, used for the
characterization of the microstructure
3.2.3
Hardness Test
The Vickers hardness tests were performed using Matsuzawa DVK-2 material
testing machine, at five different locations of the samples according to ASTM-E98-82
standard, Figure 3.5
36
(a)
(b)
Figure 3.5: (a) Matsuzawa DVK-2 used for the hardness testing (b) location of the test
on the sample
37
3.2.4
Tensile Test:
The tests were conducted according to ASTM-A370 on an instron universal
testing machine with a tensile speed of 1 mm/min. The test sample has a gauge length of
25mm and thickness of 10mm. Extensometer was used to measure the elongation. Data
presented in this report were average value of 4 separated measurements.
Figure 3.6: Instron universal tensile testing machine used in the tensile testing of the
samples
38
3.2.5
Electrochemical Measurement:
The electrochemical test ware performed using a (Parstat-2263) test set up. A
saturated calomel electrode (KCL) and graphite electrode were used as the reference and
counter electrodes, respectively. Samples (working electrode) with a cross section of 1
cm2 was used, all the polarization curves were measured at a scan rate of 0.9 mV/s.
Figure 3.7: Electrochemical test (Parstat-2263) set up used for the corrosion
measurement
39
CHAPTER 4
RESULTS AND DISCUSSION
4.1
Selection of Optimum Zinc addition
4.1.1
Nominal Composition Analysis
Chemical composition as-cast samples were determined by Energy Dispersive
Spectrometer (EDS) detector attached to Scanning Electron Microscopy (SEM), as
shown in Table 4.1.
Table 4.1: Nominal chemical composition of the Mg-Zn alloys (wt. %)
Composition
Zn
Mg
Balance
Sample 1
Sample 2
Sample 3
Sample 4
Sample 5
1.62
4.03
6.33
8.46
9.65
98.38
95.97
93.67
91.54
90.35
100
100
100
100
100
40
4.1.2
Microstructural Characterization
Optical microstructures of as cast Mg–Zn alloys were shown in figure 4.1. All
materials show nearly equiaxed grain structure. However, differences are noted among
these samples. Small and separated precipitates are observed within grains for alloys and
the width of the grain boundary becomes thicker as the content of zinc increases. In
addition, the grain boundaries are characterized by a discontinuous distribution of small
precipitates.
Figure 4.1: Microstructure of the as cast (a) pure magnesium, (b) Mg-2Zn,
(c) Mg-4Zn, (d) Mg-6Zn, (e) Mg-8Zn, (f) Mg-10Zn
41
In trying to point out what the precipitation in Figure 4.1(e) and (f) constitutes of, Figure
4.2(a) and (b) show the FE-SEM micrographs of Mg-8Zn, and Mg-10Zn alloys
respectively, and the EDS analysis conducted on the precipitation along the grain and
the grain boundary indicates that it is rich in zinc and small amount of magnesium
Figure 4.3(a) and (b) suggesting that the precipitation consist of zinc and magnesium.
Fine second phases with size of 1-2 μm can be seen in the Mg-Zn alloy in the grain and
in the grain boundary of the alloy Figure 4.2(c) and (d).
Figure 4.2: FE-SEM micrographs of (a) Mg-8Zn alloy, (b) Mg-10Zn Alloy, (c) (Mg,
Zn)-containing phase in the grain, (d) (mg, Zn)-containing phase at the grain boundary
42
Figure 4.3: EDS analysis of the secondary phases (a) on the grain, (b) at the grain
boundary
43
4.1.3
Hardness Test
With respect to pure magnesium in figure 4.4 the Vickers hardness values of Mg
xZn (2, 4, and 6) as a function of the %wt Zn follow an increasing pattern but slightly
drop at Mg-8Zn, and Mg-10Zn alloys. This slight drop in hardness value may be
attributed the formation of the secondary phases on these alloys Mg-8Zn, and Mg-10Zn.
Because the maximum solubility of zinc in magnesium is 6.2% wt and as highlighted by
Yin Dongsong [31] the excess Zinc reacts with Mg and form large amount of Mg, Zn
containing phases in the matrix and grain boundary. These phases segregate the matrix
and increase the number of crack sources. Therefore, the strength/hardness of the alloy
will not be improved.
Figure 4.4: The hardness value of Mg-Zn alloys as a function of zinc addition.
44
4.1.4
Tensile Test
The tensile strength of the alloy with respect to %wt Zn addition in figure 4.5
also shows an improving trend until 8%wt Zn where the value drop but rise again at
10%wt Zn. The value of the tensile test should follow similar trend as the hardness
graph, but this dissimilarity may be related to error or inconsistency in the experiment or
the test machine especially for Mg-10Zn, even though the student was very careful to
ensure consistency throughout the project. On the other hand the Yield strength value in
Table 4.2 shows an ascending pattern of strength for all the alloys. But some rise and fall
were observed in the value of the elastic modulus of the alloys. These phenomena of the
may be fully explain by the formation of secondary phase as has been shown in figure
4.1(e, f) and 4.2(a, b). As described by Yin Dongsong, excess Zn results in the formation
of second phases when reacts with magnesium and become sources of separation in the
matrix and the grain boundary, hence, the tensile strength of the alloy will drop.
Tensile strenght (MPa)
Tensile
180
160
140
120
100
80
60
40
20
0
159.57
134.82
105.57
139.84
108.28
Mg-Zn alloys
2%Zn 4%Zn 6%Zn 8%Zn %10Zn
wt% zn
Figure 4.5: The Tensile strength value of Mg-Zn alloys as a function of zinc addition.
45
Table 4.2: The Tensile strength, Yield, and Elastic Modulus value of Mg-Zn alloys
Alloys
Tensile Strength
(MPa)
Yield strength
(MPa)
Modulus
(GPa)
Mg-2Zn
105.57
54.19
35.94
Mg-4Zn
134.82
65.61
29.93
Mg-6Zn
139.84
83.43
48.66
Mg-8Zn
108.28
89.98
33.29
Mg-10Zn
159.57
128.80
41.55
46
4.1.5
Corrosion Electrochemical Test
Figure 4.6 shows the electrochemical polarization curves of the Mg-xZn alloys
(x = 2 - 10wt %). It could be seen that the cathodic polarization current of hydrogen
evolution reaction (-1.675V) on Mg-6Zn alloy sample was greater than that on Mg-Zn
(2, 4, 8, and 10) alloy samples, which indicated that the over potential of the cathodic
hydrogen evolution reaction was lower for Mg-6Zn alloy sample. Meaning to say Mg6Zn is less prone to corrosion compare to the other samples. According to Zijian Li [28]
generally, the cathodic polarization curves were assumed to represent the cathodic
hydrogen evolution through water reduction.
Figure 4.6: Electrochemical polarization curves of Mg-Zn alloys under investigation
47
CHAPTER 5
CONCLUSION
Developing optimum Mg-Zn binary alloy composition in terms of tracing the
effect of zinc addition in magnesium was the major objective of this project. So,
therefore it can conclude that:
i.
Mg-6Zn alloy with the hardness value of 74.44HV, tensile strength of
139.88MPa, and modulus of 48.66GPa could be considered the optimum
composition based on this project work. It shows significant improvement in
respond to the zinc addition, virtually better than the other composition studied.
Likewise in terms of resistance to corrosion it shows higher potential then the
rest of the alloys.
ii.
In comparison with the reported hardness and tensile strength of the hardest
natural bone (16-168HV, and 50-150), and due to the fact that implant material
should not be harder than the natural bone, it can be concluded that the
properties exhibits by Mg-6Zn are comparable.
48
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51
APPENDIX A
Compositional Analysis of the as- cast Samples
Compositional Analysis of the as- cast Samples of Mg-2Zn
52
Compositional Analysis for the as- cast Sample Mg-2Zn
53
Compositional Analysis for the as- cast Sample Mg-4Zn
54
Compositional Analysis for the as- cast Sample Mg-4Zn
55
Compositional Analysis for the as- cast Sample Mg-6Zn
56
Compositional Analysis for the as- cast Sample Mg-6Zn
57
Compositional Analysis for the as- cast Sample Mg-8Zn
58
Compositional Analysis for the as- cast Sample Mg-8Zn
59
Compositional Analysis for the as- cast Sample Mg-10Zn
60
Compositional Analysis for the as- cast Sample Mg-10Zn
61
APPENDIX B
The Stress/Strain Graphs for the alloys Samples
The Stress/Strain Graphs for Mg-2Zn
62
The Stress/Strain Graphs for Mg-4Zn
63
The Stress/Strain Graphs for Mg-6Zn
64
The Stress/Strain Graphs for Mg-8Zn
65
The Stress/Strain Graphs for Mg-10Zn
66
APPENDIX C
Polarization Curves of the Samples
Polarization Curves of Mg-2Zn
Mg-2Zn
67
Polarization Curves for Mg-4Zn
Mg-4Zn
68
Polarization Curves for Mg-6Zn
Mg-6Zn
69
Polarization Curves for Mg-8Zn
Mg-8Zn
70
Polarization Curves for Mg-10Zn
Mg-10Zn
71