In vitro degradation behaviour and fracture mechanism of Mg-0.8Ca under a constant strain

Nur Amirah binti Shaharom Supervised by: i) Dr. Hendra Hermawan ii) Assoc. Prof. Ir. Dr. Mohammed Rafiq bin Abdul Kadir

Abstract

Magnesium-based alloys are currently attracted much interest owing to its

potential for biodegradable implants application. However, despite many of its

advantages which are superior to those of current implant materials, tackling the

rate of degradation of magnesium alloys is one of the grand challenges

highlighted. Mg-0.8Ca is believed to have optimum qualities to serve as implant

material, thus have been chosen as the testing materials in present work. Failure

mechanism of tested material is proposed to be evaluated to further understand

its degradation behaviour under synergistic presence of mechanical loading

along with corrosive physiological environment. It is believed that under such

condition the materials are highly susceptible to stress corrosion cracking (SCC)

which might lead to catastrophic complications and sudden failure. This study

aims to evaluate the susceptibility Mg-0.8Ca towards SCC in physiological

environment, determine the mode of SCC crack propagation and recognise the

causes of failure which will be depicted from the propose testing methods. The

understanding of failure mechanism will give contributions in the improvement of

material properties in order to be successfully applied as biodegradable implants.

Introduction

Metallic biomaterials are predominately applied for impaired bone tissue fixture and

replacement [1] due to comparable monotonic strength they exhibit for such purposes.

However, evaluation on current metallic materials used [1-5], i.e. stainless steel,

titanium and cobalt-chromium-based alloys, revealed a deleterious complications due to

debris produced from loading-related scenarios, in combination with corrosive

environment of the implantation site. Stiffness ratio of current metallic implants toward

the adjacent bones is relatively huge leading to various clinical issues due to stress

shielding effects [5, 6]. Table 1 below shows typical properties of current metallic

materials in comparison to those of human bone.

Table 1 Mechanical properties of commonly used metallic biomaterials and human bone [4]

Stainless

steel Cobalt-chromium

alloy Titanium

alloy Magnesium

Human bone

Density (g cm-3

) 7.9 8.1 8.3 9.2 4.4 4.5 1.74 2 1.8 2.1

Youngs modulus (GPa)

189 205 230 110 117 41 45 3 20

Compressive yield strength (MPa)

170 310 450 1000 758 115 65 100 130 180

Development of orthopedic implants had recently shifted from classic standard of

material selection to a rather socially and economically beneficial. The selection is now

demanded the materials to be biologically active in enhancing physiological response

and implant-tissue interactions while progressively degrade during the healing period of

damaged tissue [5]. Designated to avoid secondary surgical procedure for implant

removal, biodegradable magnesium-based implants which generate mainly soluble,

non-toxic degradation products have been generally approved to serve as bone fixation

devices [2, 5, 7]. With sufficient mechanical strength and load-bearing capacity,

magnesium-based implants exceeded cotemporary polymer based implants for

orthopaedic applications. Furthermore, Youngs modulus of magnesium-based implants

is a lot closer to human bone compared to permanent metallic implants currently used

[4] which may contribute to the reduction of stress shielding effects [5]. However, the

rate of degradation of magnesium-based alloys is in biggest concern [1, 2, 4, 5, 8, 9] as

the degradation process evolves intolerant hydrogen gas accumulation and the implant

might lose its mechanical integrity before the fractured bones completely heal [4, 10].

Generally, poor corrosion resistance in magnesium alloys is believed to be influenced

by a few factors which include [11] (a) corrosion product films are too weak to inhibit the

high intrinsic dissolution tendency of magnesium, and (b) speedup of micro-galvanic

localized corrosion due to the presence of second phases serving as local cathodes. It

is believed that understanding degradation mechanism of magnesium alloys is the best

perspective in establishing a material with optimised properties most suited the implant

applications.

Alloying is one of the methods currently developed to enhance magnesium mechanical

properties and corrosion resistance [2, 4, 5, 7, 11-16]. Addition of alloying elements has

been observed analytically producing magnesium alloys with a lower corrosion rate than

that of high purity magnesium in 3% NaCl [11]. Among various alloying elements

available, calcium has been favourably chosen and considered as a unique alloying

addition owing to Mg2Ca phase features exhibited in low alloyed Mg-Ca systems which

further affect its microstructure, mechanical properties, electrochemical behaviour and

degradation kinetics [5, 7, 14, 15]. Addition of calcium in Mg-Ca alloys reduces grain

size and improves general and pitting corrosion resistance as well [5, 14]. Li et al. [14]

highlighted a few considerations in choosing calcium as alloying element in magnesium

for biodegradable material which include the composition of calcium in human body, its

close density to those of human bones, and the need of magnesium in calcium

incorporation into the bone have been reported [17] with potential benefit for bone

healing. Corrosion in Mg-Ca alloys with calcium content up to 0.8 wt% is well distributed

and homogenous in texture while those with higher calcium content demonstrate

uncertain, dissipated corrosion. Magnesium with 0.6 0.8 wt% calcium has been

proven exhibited slowest degradation rate [9, 15] and good biocompatibility [7].

Figure 1 Hydrogen evolution of Mg-Y in 0.1M NaCl. The rate of degradation can be represented by instantaneous

value from hydrogen evolution curve or by average value over a particular time period [11].

Most assessments done in vitro considered only on general practice of magnesium

alloys for orthopaedic implants, instead of simulating the implantation site condition

specifically. Improvement in corrosion resistance and mechanical properties focused on

the effect of alloying element and mechanical processes only, and despite of failure

reported experienced by current metallic biomaterial devices, no further assessment on

the mechanism of failure in corrosive environment for biodegradable implants has been

made despite of the fact that SCC in magnesium alloys would give severe effects to the

materials when in service. Thus, the present work utilizes every aspects of Mg-0.8Ca

which have been evaluated as the absolute choice of material to be applied specifically

for craniofacial fixation devices. Since the magnesium alloys corrode very fast in

chloride environment, the occurrence of SCC needs to be ascertained whether the loss

of mechanical properties is resulted from the SCC or just the reduction in specimen

cross-section due to anodic dissolution [18]. Understanding fracture mechanism

endures by Mg-0.8Ca under constant strain in simulated physiological media is

expected to contribute further improvement on its corrosion resistance and service life.

Literature Reviews

Current developments of metallic materials encompass groups of materials with fine

interaction towards biological environment and capable to progressively degrade during

the healing process of damaged tissue [5]. Instead of relying upon polymers to be

applied as implant materials, biodegradable metallic substitutes are worthier to be

established on account of better load-bearing capacity and mechanical strength they

demonstrate [5]. Magnesium and magnesium-based alloys are currently drawing much

interest as promising candidates to be applicable for such purposes. Current research

development on magnesium-based biodegradable implants is focusing mostly on Mg-Al

and Mg-RE alloy systems [14]. In vitro assessment [19] of degradation rate and

mechanism for magnesium has been summarized by Kirkland et al. detailing their

advantages and limitations.

Witte et al. [20] have conducted an investigation to compare the corrosion rates of

AZ91D and LAE442 and concluded that corrosion rates measured in vivo were orders

of magnitude lower than those measured in substitute ocean water prepared according

to ASTM-D1141-98 [21]. In vitro test has been accomplished using various solutions

which include; (i) simulated body fluid (SBF), (ii) artificial plasma (AP), (iii) phosphate

buffer saline (PBS), and (iv) minimum essential medium (MEM, Invitrogen). In vitro test

[16] conducted to evaluate the corrosion rate of various magnesium alloys reported that

high purity magnesium had the lowest corrosion rate which increased with increasing

amount and effectiveness of second phase. Song [16] concluded that Zn-containing

magnesium alloy is the most promising material to be developed into a biodegradable

material with some modifications such as anodised coating to delay the degradation and

purification to reduce the rate.

A study [22] conducted to understand failure mechanisms of commercial Kntscher

nails of type 316L stainless steel under the combined action of corrosion and stress has

been done by G. Bombara and M. Cavallini. Evaluation on damage morphology of the

device showed transgranular, branched and unblunted cracks which were first identified

as SCC behaviour rather than corrosion fatigue (CF). This was corroborated with

appearances of minor and indeterminate fretting fractographs using scanning electron

microscope (SEM). The results obtained agreed with investigation done by M.

Sivakumar and S. Rajeswari [23] which concluded SCC of failed intramedullary nail has

been worsened by high inclusion content and large grain size of the implant.

Fractography of failed commercial AISI stainless-steel 316L presented by S.

Bandyopadhyay and P. Brockhurst [24] has shown evidence of SCC failure with crack

propagation both in intergranular and transgranular modes which determined the

susceptibility of austenitic stainless steel at a rather low temperature than mostly

reported.

Kannan and Raman [25] had claimed that SCC susceptibility of Mg-Al-Zn alloy is not

substantial and should not be a concern for implant applications despite of stress-

assisted failures that have been reported [2, 12, 26, 27] for current orthopaedic

implants. However, a critical review done by Atrens et al. [11] highlighted a number of

grand challenges in development of magnesium-based implants, which include

consideration of SCC due to corrosion-assisted stresses that is known existed in

implant-bone systems. This review is in agreement with investigation by Choudhary and

Raman [18] which had confirmed the SCC susceptibility of both smooth and notched

specimens of a magnesium alloy in physiological environment. In a different

investigation that has been done by Choudhary et al. [28] on AZ91D in modified SBF

and air at different strain rates revealed that the alloy is susceptible to SCC.

Materials and Methods

Sample preparation

Samples will be used are cast Mg-0.8Ca samples whose geometry and dimensions

presented in Figure 2. The sample will be ground successively to 1200 grit using SiC

paper, ultrasonically cleaned in acetone, and dried to room temperature for 24 hours

prior to testing.

Figure 1 Dimension of the proposed sample.

Microstructural characterisation

The samples which will be used for microstructure characterisation will be mechanically

ground with SiC paper up to 2500 grit and polished, follow by chemical etching in 2%

Nital solution. Surface morphology of the samples before and after the introduction of

strain will be observed using optical microscope. Furthermore, microstructure of

samples surface will be examined using SEM before and after the samples failure.

Failed samples will be first cleaned chemically using 20% CrO3 and 10% AgNO3

solution. The changes of microstructure morphology will be observed to further

understand the influence of strain towards failure of samples and is believed to be

closely related to the fracture mechanism. The characterisation of failed samples is

expected to have features of brittle and/or ductile fracture, and show the mode of

cracking either transgranular or intergranular.

Immersion test

The samples will be assembled to a custom-made apparatus for immersion test in

simulated body fluid (SBF) with pH value will be adjusted to 7.4 and strain variation of

0.25% and 0.5% elongation. The test will be done for 1, 2, 4, and 6 hours to evaluate

the rate of degradation under such condition from the weight loss percentage, and then

prolonged until failure. Each test will be performed in triplets in order to examine the

reproducibility. Another set of immersion test will be conducted as control experiment.

The samples for the control experiment will not be put under strain, but evaluate the

same measurement of weight loss for the same time intervals, which will then be

prolonged until failure. All samples will be taken out, left to dry at room temperature for

24 hours and be measured of the weight loss, Wa before chemical cleaning. Then, the

samples will be cleaned with chromic acid solution to remove corrosion products and let

dry to room temperature before the final weight loss, Wb will be recorded.

Wb from failed samples from immersion test with introduction of strain will be compared

to Wb of the samples from immersion test without strain at the same failure time. The

microstructure of both samples will be examined and is expected to verify the effects of

strain on the samples microstructure and their influences on degradation rate.

In situ observation

A sample from the immersion test with strain value of 0.5% will be randomly picked for

in situ observation of failure using stereo microscope. Fractography and failure

mechanism are expected to be better visualised using this method.

Expected results and discussion

The degradation rate will be computed from experiments conducted in different

conditions; (i) sample will be put under strain in SBF, (ii) sample will be put under strain

in air, and (iii) sample will be immersed in SBF for 24 hours then be put under strain in

air. Each of the experiments however, will be repeated thrice. The variation in testing

condition is to recognise the causes of failure [18]; either it is indeed a result of the

synergistic effect of stress and corrosive environment (i.e. SCC) or simply attributed to

the continuously reducing cross-sectional area of the sample due to high corrosion rate

of magnesium alloys.

Table 2 Design of experiment that is expected to be achieved with some data obtained from current conducted

experiments.

Media Displacement Time to failure Weight Loss pH Degradation Rate

SBF

EXPERIMENT 1 - Degradation Rate of MgCa Under Tension Stress

0.50%

Time (min) W W W - W Wb W - Wb pH pHi (W - Wb)/t

60 1.87 1.869 0.001 1.838 0.032 7.2 7.9 0.000533333

120 1.89 1.886 0.004 1.852 0.038 7.2 7.9 0.000316667

240 1.929 1.927 0.002 1.835 0.094 7.2 8.2 0.000391667

360 1.803 1.804 -0.001 1.656 0.147 7.4 8.4 0.000408333

tf - 960 1.94 1.86 0.08 1.654 0.286 7.3 9.2 0.000297917

EXPERIMENT 2 - Degradation Rate of MgCa Under Tension Stress

0.25%

Time (s) W W W - W Wb W - Wb pH pHi (W - Wb)/t

60 1.75 1.78 -0.03 1.655 0.095 7.1 7.6 0.001583333

120 1.78 1.82 -0.04 1.727 0.053 7.1 7.6 0.000441667

240 1.88 1.9 -0.02 1.788 0.092 7.1 7.8 0.000383333

360 1.75 1.73 0.02 1.64 0.11 7.1 7.9 0.000305556

T2 W20

EXPERIMENT 3 - Degradation Rate of MgCa Under Tension Stress

0%

W W W - W Wb W - Wb pH pHi

1h W21

2h W22

4h W23

6h W24

T5 W25

Results obtained in Table 3 are from the first round of experiments. From the data,

some graphical evaluation could be made as shows in Table 3. Computing stress vs.

time curves will show changes in mechanical properties of the sample tested for each

assessed conditions. It is expected that maximum reduction in mechanical properties

which might fail the fastest among the samples tested is the one under strain in SBF. It

is expected that stress-time curve for the samples that will be pulled in air indicate some

similar behaviours in comparison to those pre-immersed samples.

Table 3 Degradation rate evaluated for current conducted experiments.

Fractograph of the sample that will be tested under strain in SBF is expected to show

common mode for SCC crack propagation of magnesium alloys, i.e. localised attack at

the sample circumference which in higher magnification, will reveal transgranular

cracking and localized cracks (Figure 3).

60 120 240 360 tf - 960

0.50% 0.00053333 0.00031667 0.00039167 0.00040833 0.00029792

0.25% 0.00158333 0.00044167 0.00038333 0.00030556

0

0.0002

0.0004

0.0006

0.0008

0.001

0.0012

0.0014

0.0016

0.0018

Weig

ht

Lo

ss (

g)

Rate of Degradation (gm-1)

Degradation rates of Mg-0.8Ca under various strain values in SBF

Figure 2 Factograph of failed sample with common mode of SCC cracks propagation for magnesium alloys: (a)

overall fracture surface, and (b) evidence of trransgranular cracking [18].

The experiment will use weight loss method in evaluation of degradation rate of Mg-

0.8Ca sample. At least two from three types of measurement, i.e. weight loss rate,

hydrogen evolution rate and rate of Mg2+ leaving the metal surface, are recommended

by Atrens et al. [11] to be implemented to better clarify the rate of corrosion on

magnesium alloy in service. However, since this study is not focus on the rate of

corrosion, only weight loss rate will be measured to give a framework of service and

learn the impact of such simulated condition on the specimen towards fracture. It is

believed that the rate of corrosion is closely related to the failure mechanism of

implants.

There are two types of apparatus for transgranular SCC evaluation, the Linearly

Increasing Stress Test (LIST) and the Constant Extension Rate Test (CERT). Review

done by Atrens et al. [11] described that both of these methods are useful to identify the

occurrences of SCC and have capabilities to measure the threshold stress and crack

velocity from the final crack size divided by the time of cracking when coupled in

identification of crack initiation. The increased crack propagation time under CERT

conditions may be important in determining the mechanism for SCC in magnesium

alloys. However, the experimental setup presented in this study will be custom-made,

an apparatus that simulate the condition which initiate conventional SCC. There may be

a room for improvement by applying either LIST or CERT, or both to better evaluate

failure mechanism of proposed material.

Choudhary and Raman [18] had done SSRT under different testing conditions which

included continuous cathodic charging and simultaneous strain in SBF for indication of

possible combined effect of hydrogen-induced cracking and anodic dissolution.

Furthermore, circumferential notch tensile (CNT) testing was used in their work to

determine the threshold stress intensity factor for SCC (KISCC) of the tested samples in

SBF. SEM fractographs (Figure 4) of the failed samples from CNT test had provided

better description on the failure mechanism.

Figure 3 SEM fractographs obtained from CNT: (a) the overall fracture surface showing machined notch, fatigue pre-

crack, SCC and mechanical failure zones, (b) fatigue pre-cracked and SCC zones (arrow indicate crack propagation

direction), and (c) mechanical failure zone showing dimples [18].

Intergranular in magnesium alloys occurs typically by micro-galvanic corrosion, when

there is a presence of continuous or nearly continuous second phase along the grain

boundaries, a microstructure typical of many cast creep-resistant magnesium alloys. In

contrast, transgranular SCC occurs through the -phase magnesium matrix. From the

review done by Atrens et al. [11], they concluded that SCC is associated with

environmental conditions leading to the local breakdown of a partially protective surface

film which is usually caused by localized corrosion by chloride ions. However, load may

be influenced the breakdown as well since the occurrence of SCC is observed on pure

magnesium and magnesium alloys (AZ91, AM60, AS41, ZK60A-T5) in distilled water.

The influenced of hydrogen to the fracture behaviour and the effect of alloying elements

on SCC propagation are not well understood currently. In advance, the understanding of

microstructure influences on SCC is required to better understand magnesium

transgranular SCC. Furthermore, the understanding of effective hydrogen diffusion rate

to the fracture process zone needs to be included which requires an understanding of

hydrogen-trap interactions in magnesium alloys. Although Kannan and Raman [13]

found insignificant decreases of strength and ductility for AZ91 in simulated body fluid

using CERT which indicate SCC, such failure is still largely unexplored.

Conclusion

Current studies on the possibilities to implement magnesium-based alloys as

degradable implants demanding a much deeper understanding of the degradation

behaviour, particularly involving microstructural analysis, to tackle rapid corrosion of the

material in physiological environment. SCC is one of many failures recognised among

metallic materials, thus it is demanded to have absolute evaluation on such issues to

provide crucial information on whether the materials will be tolerable for the given

service time of the device in practical. Such studies have been done on industrial

magnesium alloys, which are less likely applicable as implant materials. Thus, the

present work emphasised the SCC susceptibility assessment and the mechanism of

SCC involved, if applicable, leading to failure of the proposed implants material, Mg-

0.8Ca.

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