REVIEW

Biodegradable magnesium implants for orthopedic applications

Hazibullah Waizy • Jan-Marten Seitz • Janin Reifenrath •

Andreas Weizbauer • Friedrich-Wilhelm Bach • Andrea Meyer-Lindenberg •

Berend Denkena • Henning Windhagen

Received: 9 March 2012 / Accepted: 9 May 2012

� Springer Science+Business Media, LLC 2012

Abstract The clinical application of degradable ortho-

pedic magnesium implants is a tangible vision in medical

science. This interdisciplinary review discusses many dif-

ferent aspects of magnesium alloys comprising the manu-

facturing process and the latest research. We present the

challenges of the manufacturing process of magnesium

implants with the risk of contamination with impurities and

its effect on corrosion. Furthermore, this paper provides a

summary of the current examination methods used in in

vitro and in vivo research of magnesium alloys. The

influence of various parameters (most importantly the

effect of the corrosive media) in in vitro studies and an

overview about the current in vivo research is given.

Introduction

The quality and speed of bone healing depends on the size

of the fracture gap and the achieved stability [1]. Therefore,

an adjusted stability is essential for fracture healing. Early

recovery after bony healing is necessary to prevent long-

term disabilities and hospitalization periods, as well as the

risk of malunion, nonunion, and infection. Intramedullary

nailing and the usage of pins, screws, or plates are accepted

approaches for fracture stabilization. Stainless steel and

titanium are the most frequently used materials in medical

applications in the last few decades. These currently

applied permanent metallic internal fixation devices have

several negative aspects such as stress shielding [2, 3], an

inflammatory osteolysis caused by released toxic titanium

particles [4], interference in radiological studies [5], and

the need of a second surgery for implant removal.

In recent years, biodegradable biomaterials for medical

use have gained interest and are intensively investigated;

promising candidates are magnesium alloys. These bio-

materials have to comply several requirements: (i) good

biocompatibility and non-toxicity of degradation products,

(ii) appropriate mechanical properties, and (iii) a moderate

degradation rate adapted to the fracture healing process.

Mg2? is an essential component of the human body and it

is mainly stored in bones. Many enzymes require magne-

sium as a co-factor to provoke a chemical reaction for

example DNA replication [6]. The recommended daily

intake of magnesium is approximately 300 mg for adults

[6]. Human adult blood plasma obtains a total Mg con-

centration of 0.65–1.05 mmol/l [7].

In addition, magnesium possesses desirable mechanical

properties (density: 1.74 g/cm3; elastic modulus 45 GPa,

and compressive yield strength 65–100 MPa) closer to

those of natural bone than currently used titanium alloys

H. Waizy (&) � A. Weizbauer � H. Windhagen

Department of Orthopaedic Surgery, Hannover Medical School,

Anna-von-Borries-Str.1-7, 30625 Hannover, Germany

e-mail: Hazibullah.Waizy@ddh-gruppe.de

J.-M. Seitz � F.-W. Bach

Institute of Materials Science, Leibniz University of Hannover,

An der Universitat 2, 30823 Garbsen, Germany

J. Reifenrath

Small Animal Clinic, University of Veterinary Medicine

Hannover, Bunteweg 9, 30559 Hannover, Germany

A. Meyer-Lindenberg

Clinic for Small Animal Surgery and Reproduction, Centre of

Clinical Veterinary Medicine, Faculty of Veterinary Medicine

Ludwig-Maximilians-Universitat Munchen, Veterinarstr. 13,

80539 Munich, Germany

B. Denkena

Institute of Production Engineering and Machine Tools (IFW),

Leibnitz University of Hannover, An der Universitat 2, 30823

Garbsen, Germany

123

J Mater Sci

DOI 10.1007/s10853-012-6572-2

[4, 8]. These mechanical properties of magnesium mini-

mize the disturbance of bone growth and remodeling by

reduced stimulation (‘‘stress-shielding’’) [2]. An approach

of enhancing the mechanical properties for clinical appli-

cations is alloying [9]. Most commonly used for alloying

are elements like calcium, lithium, zinc, zirconium, rare

earth metals, or aluminum [10]. Some of these elements,

most notably aluminum and rare earth metals, are sus-

pected to cause adverse effects in organism [11, 12].

In physiological fluids, magnesium and its alloys

degrade according to the following reactions [13]:

Anodic reaction : Mg ! Mg2þ þ 2e�

Cathodic reaction : 2 H2O þ 2e� ! H2 " þ 2 OH�

Mg2þ þ 2 OH� ! Mg OHð Þ2 sð Þ

The corrosion of magnesium and its alloys are accompa-

nied by hydrogen-gas development. The production of high

amounts of gas in a short period of time is not desired for

clinical application. The gas development is dependent on

the corrosion rate. One of the first clinical applications of

pure magnesium implants was in 1906 by Lambotte [14].

The results of Lambotte at this early stage of magnesium

research were not satisfactory and the pure magnesium

plates were removed due to extensive gas cavities, local

swelling, and heavy pain [14].

In this review, we provide a survey about many different

aspects of recent magnesium alloys research and the dif-

ficulties in the magnesium manufacturing process by con-

tamination primarily with iron, nickel, and copper and its

effect on the corrosion rate. A series of different in vitro

studies investigated the degradation behavior of magne-

sium alloys. The influence of various parameters (most

importantly the choice of the corrosive media) is presented

in that paragraph. In vivo evaluation of magnesium alloy

samples plays an important role and the current state of

research is given.

Challenges in magnesium alloying: the risk

of contamination with impurities

Although pure magnesium demonstrates generally suitable

corrosion properties as an implant material for resorbable

applications, it frequently possesses insufficient mechani-

cal properties. One possibility of enhancing its mechanical

properties is represented by the use of magnesium alloys.

By means of alloying with suitable alloying elements, it is,

for instance, possible to eliminate the mechanical defi-

ciencies of pure magnesium. Here, however, it must be

taken into consideration that such elements simultaneously

modify the material’s corrosion behavior (rate and type of

corrosion) [15].

As a rule, the effects which the alloying elements pro-

duce in this respect are deliberate and intentionally

obtained. However, such effects can also be unintentionally

generated by means of contaminants during the implants

manufacture. Elements such as iron, nickel, and copper,

which are recurrently found in magnesium alloys are fre-

quently unintentional and due to the manufacturing process

[16]. In this respect, iron in particular plays a significant

role owing to its frequently inevitable contact with mag-

nesium alloys during their manufacture. During the melting

and stirring of the melt in the steel crucible, which occa-

sionally contains nickel, the iron and nickel contents can

precipitate out of the crucible’s wall and thus end up in the

magnesium’s melt in which it represents just one constit-

uent of the alloy [17, 18]. Contaminants containing copper

mainly arise by employing ‘‘impure’’ aluminum alloying

elements [19]. The solubility characteristics of the con-

taminating elements in magnesium are as follows:

The equilibrium diagram (Fig. 1) for magnesium and

iron shows a eutectic at 650 �C which lies close to the

melting point of magnesium. The solubility of iron in

magnesium is limited to 0.008 at.% within this eutectic.

Above this, the element exists as alpha-iron in the mag-

nesium matrix [20]. Typical melting or, as the case may be,

casting temperatures for alloying magnesium lie in the

range of 650 and 700 �C depending on the added alloying

elements. Table 1 depicts the fraction of dissolved iron

from various sources at precisely these temperatures. The

fractions range from 0.0044 to 0.0218 at.%. The solubility

of nickel in magnesium is reported to be very small and lies

significantly below 0.04 at.% at a temperature of 500 �C

[21, 22]. Likewise, the solubility of copper in magnesium is

presented as extremely small. Here, the solubility is

between a range of 0.153 and 0.191 at.% at a temperature

of 485 �C [23].

Fig. 1 Curve to illustrate the tolerance limit for a contamination of

magnesium with the elements Fe, Ni, Cu, according to [15]

J Mater Sci

123

Even in small amounts, iron, copper, and nickel as

constituents in pure magnesium lead to an increase in the

corrosion rate. It was possible for Ref. [15, 24, 25] to

demonstrate this in comparison with other elements (Ag,

Ca, Zn, Cd, Sn, Pb, Al,…). The corrosion rates of binary

alloys were examined in (3 % NaCl) saline solution. In

doing this, concentrations of iron, copper, and nickel,

smaller than 0.2 %, lead to a significant increase in the rate

of corrosion [15, 26]. This increase of the corrosion rate is

attributed to the elements’ low solubility in magnesium as

well as their distinctly more noble position in the electro-

chemical series [15, 27, 28]. A particularly large impact on

the corrosion rate of magnesium and its alloys is accepted

as a contamination by iron, if nothing else, because of its

frequent occurrence. Here, a galvanic couple is formed in

the magnesium’s microstructure in which the undissolved

iron particles within the magnesium matrix represents the

cathodic pole and reduces the neighboring magnesium

phases [29, 30]. Owing to the low solubility of nickel in

magnesium, the effect on the corrosion rate is described in

even more drastic terms since elementary, galvanic cells

are already formed early without the ability of compen-

sating for these cells in the form of phases [31, 32].

Although copper contaminants also demonstrate a signifi-

cant influence on accelerating corrosion, it must, however,

exist in higher doses owing to its good solubility in mag-

nesium [33].

To be able to define tolerable amounts of contaminating

elements in magnesium, the concept of tolerance limits is

introduced. According to this, the tolerance limit is given at

a location at which the contaminant’s concentration leads

to a significant acceleration of corrosion (Fig. 1). Before

reaching the tolerance limit due to contaminating elements,

the corrosion rate appears low and thereby tolerable [15].

In purely binary alloys of as-cast magnesium and iron,

nickel, or copper, the tolerance limits result as depicted in

Table 2. By alloying with a third element, the measured

tolerances can, in part, be considerably shifted. As an

example with respect to this, on combining magnesium,

aluminum, and iron, a Fe–Al phase is formed which

reduces the mitigation of tolerable amounts of iron in the

magnesium alloy to 5 ppm (for 7 % aluminum). The rea-

son for this is the presence of the FeAl3 phase which is

more electrochemically unfavorable for magnesium alloys

and, in comparison with pure iron particles, acts more

nobly and therefore more corrosion intensively in magne-

sium [34, 35].

The corrosion promoting properties of iron, nickel, and

copper contaminants within magnesium alloys require

counteractive measures to enable the material’s properties

to be consistent and reproducible. In order to prevent the

introduction of iron and nickel quantities during manu-

facturing by casting technology, one can dispense with

materials for the crucible, stirrer, and dies which contain

such elements. Here, materials based on titanium represent

an already real but expensive alternative. On the other

hand, it is also possible to exploit the properties of addi-

tional alloying elements. One such alloying element is

manganese which can effectively reduce the effect of the

iron content on the corrosion of magnesium materials [16].

At a higher temperature than that at which casting is per-

formed, quantities of manganese are added to the melt.

During the cooling phase down to the casting temperature,

intermetallic phases now form from the added manganese.

The iron contaminants which, owing to their density, are

deposited onto the crucible’s base and the iron concentra-

tion in the melt are lowered [36, 37]. Moreover, it is

assumed that the added manganese envelops the iron par-

ticles and thus deprives them of direct contact with mag-

nesium. In this way, only galvanic cells arise between the

manganese and the magnesium which, owing to the dif-

ference of the chemical standard potential, are less corro-

sion intensive as cells consisting of magnesium and iron

[18, 37, 38]. Using the same method, constituents of nickel

can also be bound and eliminated from the cast. However,

this is not considered as an effective procedure for manu-

facturing a highly pure magnesium alloy. In a series of

investigations using AZ alloys, even the smallest amounts

(\1 %) of manganese could lead to a significant

improvement in corrosion behavior by which they raised

the tolerance limits for both iron (20 wt.ppm) and nickel

contaminants [18, 31, 38, 39]. It was possible to relate the

corrosion behavior of magnesium alloys and, in particular,

Table 1 Solubility limits of Fe

in liquid Mg [20, 122–126]Temperature

(�C)

According

to [123]

According

to [125]

According

to [122]

According

to [126]

According

to [124]

650 0.0113 0.007 0.0044 0.0054 0.0148

–0.0087

700 0.0152 0.0157 0.0087 0.0109 0.0218

–0.0174

Table 2 Tolerance limits for as-cast magnesium in binary compo-

sition with Fe, Ni and Cu [15, 18]

Fe (ppm) Ni (ppm) Cu (ppm)

Pure Mg 170 5–10 1000–1300

J Mater Sci

123

the Fe/Mn ratio. In general, for large values of Fe/Mn

ratios, a high corrosion potential could be verified,

whereas, for small values, a slow corrosion behavior

appeared uniformly over the surface [30, 40]. It was pos-

sible to identify the relationship between the Fe/Mn ratio

and the tolerance limit for iron by means of an AZ91 alloy

[40]. Following the evaluation of the salt spray test, a linear

relationship was clearly demonstrated between the iron/

manganese ratios and the resulting tolerance limit for iron.

However, the most effective method to be able to control

the unfavorable influence of contaminants of iron and

particularly of nickel and copper currently consists of

resorting to highly pure alloying additives [16].

Owing to its numerous alloying elements and their com-

paratively high quantities, the production of the LAE442

alloy harbors a huge risk of becoming contaminated by

nickel, copper, and iron. As described above, it is possible

that additional amounts of copper are introduced into the

alloy via the alloying element aluminum [19]. Moreover, the

alloying element, aluminum, lowers the tolerance limit of

magnesium alloys for iron to about 20 wt% with which it

forms a strongly cathodic acting Al–Fe compound [29, 31].

The largest risk of contaminating with iron results from the

casting process since steel crucibles, steel stirrers, and steel

dies are conventionally employed for producing the non-

commercially available LAE442 alloy.

Surface treatments and subsequent processes like hot

extruding and rolling in general show improving impacts

on the corrosion performances of Mg and Mg alloys [41–

44]. Surface treatments like protective films and coatings

(e.g., resorbable polymers, resorbable hydroxyapatites, and

epoxy resins) prevent the substrate magnesium from direct

exposition to surrounding fluids (electrolytes) [15, 42, 43,

45–47]. The protective surface films improve the electro-

chemical behavior by surface passivation which should,

even in case the impurity tolerance limits of the substrate

material are exceeded, delay corrosion. Corrosion mecha-

nisms which are dependant on an electrolyte should be

permanently or temporarily (depending on the coating)

inhibited, while contact corrosion which exists due to the

different phases still remains [15].

Another alternative to optimize the corrosion properties

of Mg and its alloys is to use altered casting processes [48]

and/or subsequent processes like extruding [41] or rolling

[44]. Here, grain refinement occurs having a significant

impact on the metal’s corrosion properties [41, 44, 48]. In

general, grain boundaries act as physical corrosion barriers;

a smaller grain size increases the amount of grain bound-

aries which furthermore decreases the rate of corrosion

[49]. However, it remains unclear if finer grains positively

affect corrosion which occurs due to impurities. In this

case, phases of impurity contact a higher number of adja-

cent Mg grains inflicting contact corrosion on them.

In vitro evaluation of magnesium alloys

In vitro experiments provide the opportunity to examine

newly developed magnesium alloys under standardized

conditions before testing in animals. These studies are

mainly carried out (i) to test cytotoxicity or (ii) to inves-

tigate the corrosion behavior.

In vitro cytotoxicity evaluations are useful for assessment

of feasible destructive effects of magnesium alloys and its

degradation products on cell viability and proliferation.

Surveys of the different tests are given in Table 3. MTT and

XTT assays of corrosion extracts are most often performed

to quantify cell viability using metabolic markers though it is

reported that the interference between test reagent and cor-

rosion of some alloys restrict this method [10, 50]. These

assays are in accordance with EN ISO 10993-5 which

describes tests for in vitro cytotoxicity for the evaluation of

medical devices. In few studies, the cytotoxicity is evaluated

by the determination of live cells via trypan blue exclusion

method or via cell adhesion by DAPI staining [51–53]. SEM

is often applied to analyze the cell morphology on the surface

[52, 54]. The assessment of certain cell proteins and DNA

content is sometimes applied such as the expression of the

osteogenetic-specific m-RNA (collagen 1a1, alkaline phos-

phatase ALP, and osteocalcin OC) [54, 55].

Different cell lines are used to evaluate the cytotoxicity of

magnesium alloys via in vitro experiments (Table 3). The

choice of a specific cell line is important for the performance

of cytotoxicity tests [56]. A variety of cell lines are used with

different origin (animal or human). Most of the cells that are

used for the evaluation of orthopedic magnesium implants

are osteoblasts and fibroblasts. Fischer et al. [57] compared

cytotoxicity tests with cells with cancerous origin (human

osteosarcoma cell line MG63) to an isolated primary cell line

of human osteoblasts. The results showed that the tolerance

toward magnesium extracts was higher for osteoblasts in

comparison to MG63 [57]. Feyerabend et al. [56] reported

similar observations investigated in rare earth elements

(REE) on primary mesenchymal stem cells, MG63 and a

RAW 264.7 tumor-derived mouse cell line. The higher tol-

erance of the primary mesenchymal cell line toward REE-

salts was attributed to the direct impact on cancerous cell

lines [56]. It is considered to perform cytotoxicity tests with

human bone-derived cells or mesenchymal stem cells

because of their high osteogenic differentiation potential [56,

57]. These findings are in contrast to cell lines recommended

by EN ISO 10993-5.

Many in vitro studies were performed to determine the

corrosion behavior of the investigated magnesium alloys.

Electrochemical and gravimetric measurements are fre-

quently applied to study the degradation behavior.

Immersion tests are often performed to determine the

hydrogen evolution rate [10]. Song reported that the

J Mater Sci

123

evolution of 1 ml H2 corresponds to the dissolution of 1 mg

magnesium [58]. In few studies, the degradation rate was

calculated according to the weight loss method [51, 52,

59]. Sometimes the Mg2? release from the alloy was

measured by colorimetric method using xylidyl blue-I or

via inductively coupled plasma atomic emission spectros-

copy (ICP-AES) [9, 55, 60, 61]. These in vitro methods are

useful tools to determine the corrosion rate under stan-

dardized physiological conditions. A critical discussion of

the limitations of these methods and their benefit was

compiled by Kirkland et al. [62]. Many previous studies

have shown that the biocorrosion rates displayed by in vitro

and in vivo tests are difficult to correlate with faster cor-

rosion rates in vitro [63, 64]. Recent researches have shown

that this challenge in spite of optimizing in vitro experi-

ments still exists [65]. Although in vitro studies are

insufficient to reliably predict in vivo corrosion rates, these

studies are suitable for the comparison of different alloys

under standardized conditions. A survey of a variety of

commercial and experimental alloys in minimum essential

medium (MEM) at 37 �C was given by Ref. [66].

Microstructural characterizations of the surfaces are

most commonly determined by scanning electron micros-

copy (SEM) and energy-dispersive X-ray spectroscopy

(EDS) [10]. In some studies, the grain size of the alloy is

determined according to the linear intercept method

described in ASTM E112-96 [55, 61].

The influence of the corrosion medium

Many factors have an influence on the corrosion behavior,

most importantly the choice of the corrosive medium.

Previous in vitro studies were performed with NaCl solu-

tion, phosphate buffer solution (PBS), simulated body fluid

(SBF), Hank’s solution, Dulbecco’s modified eagle med-

ium (DMEM), Ringer’s solution, or MEM. The indigents

of five solutions are given in Ref. [50]. A summary of the

corrosion media used in previous works is given in

Table 4. The composition of the various corrosive media is

crucial and varies in the concentration of aggressive ions

(e.g., Cl-), buffering agents, and additives like proteins or

organic acids. Some studies investigated the influence of

the addition of proteins on the corrosion rate. Albumin is

the most abundant protein in the human blood with a

molecular weight of 66 kDa [67, 68]. The adsorption of

proteins on the surface is complex including van der Waals

forces, hydrophobic and electrostatic interactions, and

hydrogen bonding [69]. Some authors reported a decreased

corrosion rate of MgCa alloy with the addition of proteins

[51, 70, 71]. Yamamoto et al. [60] observed a decreased

Mg2? release of pure magnesium in the beginning of

immersion and a lower total release after 14 days com-

pared to solutions without protein. In agreement, this initial

corrosion protective effect of an albumin layer in SBF was

observed in other studies as well [72, 73]. The hydrogen

Table 3 Survey of cytotoxicity experiments for evaluation of magnesium alloys for orthopedic applications excluding coated samples

Test method Subject of test Cell lines

Human cells of

non cancerous

origin

Human cells

of cancerous

origin

Animal cells of

non cancerous

origin

Animal cells

of cancerous

origin

MTT assay Coloring agent MTT (3-(4,5-dimethylthiazol-2-

yl)-2,5-diphenyl tetrazolium bromide) is

metabolic transformed to formazan

Human umbilical

cord perivascular

(HUCPV) [56]

Isolation of

primary human

osteoblasts [57]

MG63 [56,

57, 73, 84]

NIH3T3 [127]

MC3T3-E1 [52,

127–129]

RAW 264.7 [56]

rabbit bone

marrow stroma

cells (rBMSC)

[9]

L-929 [101,

104, 127]

(according

to EN ISO

10993-5)

XTT assay XTT (sodium 3,3-{1-[(phenylamino)carbonyl]-

3,4-tetrazolium}-bis(4-methoxy-6-

nitro)benzene sulfonic acid hydrate) is

transformed by living cells to a

formazanproduct

L-929

(according

to EN ISO

10993-5)

Cell count and

cell

attachement

Adherent cells were stained with DAPI (4’,6-

diamidino-2-phenylindole dilactate) and

counted via fluorescence microscope,

tryptophan blue exclusion method or other cell

count methods

Bone marrow

derived

mesenchymal

stem cells [53]

SaOS2 [51] MC3T3-E1 [52,

54]

DNA content Qualitative RT-PCR of Col1a1, ALP and OC MC3T3-E1 [54]

Cell protein Determination of the alkaline phosphatse (ALP)

activity

MG63 [55] L-929 [55]

J Mater Sci

123

evolution rate of the M1A alloy was suppressed in the first

4 h; however, the long-term degradation rate was higher

with albumin [73]. Similar results were reported in the first

hours of exposure to the albumin containing medium with

WE43. The similar electrochemical behavior after 9 h in

SBF with and without protein indicated that the protein

layer is either destroyed or no longer dominant [72]. The

influence of different amount of proteins in the corrosive

solution was investigated by various studies [67, 71, 74].

Liu et al. [71] reported an accelerated corrosion inhibition

effect at higher albumin concentrations (10 g/l). In addi-

tion, Feyerabend et al. [75] reported that the Mg2? release

from the metal was increased with the addition of proteins.

Many corrosive media are buffered to avoid change of

the pH value. The choice and amount of buffer consider-

ably affects the corrosion rate. HEPES, TRIS, phosphate

buffers, and HCO3-/CO2 are widely used [68]. A study

with NaCl and PBS solution showed that PBS decreases the

corrosion rate after long term observation because of the

precipitation of phosphate containing salts on the surface.

This precipitation layer contributes to higher corrosion

resistance than magnesium hydroxide in NaCl solution

[76]. The study of Yamamoto et al. [60] investigated the

effect of the buffering reagents HEPES (N-2-hydro-

xyethylpiperazine-N0-2-ethane sulfonic acid) or NaHCO3

(sodium bicarbonate) added to NaCl on degradation

behavior. It was observed that the Mg2? release steeply

decreased in the first 2 days in NaCl ? NaHCO3 in con-

trast to NaCl and NaCl ? HEPES, which indicates the

precipitation of insoluble carbonates on the surface of

magnesium in NaCl ? NaCO3. The total amount of

released Mg2? ions after 14 days immersion was higher in

NaCl solution with HEPES than with NaHCO3. Similar

results were found in a study with different concentrations

of the buffer agent HCO3-. This study showed decreased

corrosion rate of pure magnesium with higher concentra-

tions of HCO3- in SBF. The initial degradation rate was

lower with higher buffering reagents content [77]. Contrary

results were presented in a study by Xin et al. [78] with

AZ91. The initial hydrogen evolution rate of the two

solutions with HCO3- was higher which indicated an

accelerated consumption of OH- by HCO3- which pro-

motes the dissolution reaction of magnesium. After 96 h,

the corrosion rate was more decreased in solutions with

HCO3- which was explained by protection effects of

phosphate and carbonate on the surface.

In vitro studies under cell culture conditions

and with CO2 gas application

Some studies were performed under cell culture conditions

(95 % humidity, 37 �C, 5 % CO2) [51, 53, 65, 66, 75, 79,

80]. This approach is intended to more preferably simulate

the environment in blood of in vivo studies. The main

buffering system in blood is HCO3- and dissolved CO2

[75]. Zainal Abidin et al. [81] maintained the pH value due

to constant CO2 gas application through the solution in

bubbles. In vitro studies have shown that the presence of

CO2 has a significant effect on the composition of the

corrosion product layer [75, 79]. The formation of MgCO3

and the precipitation of a crystalline structure were

observed [75, 79]. MgCO3 is suspected to possess corrosion

inhibition properties after the formation of a layer [75].

Influence of additional parameters: temperature, pH,

microstructure and flow velocity

Most of the studies were performed with a temperature in

the range of the human physiological temperature

(35.8–37.2 �C). Kannan et al. [82] reported a decreased

corrosion resistance of the AZ91 alloy at body temperature

compared to room temperature (20 ± 2 �C). Commonly,

the corrosion rate increases when the temperature is ele-

vated [15]. Magnesium and its alloy lead to an alkalization

of the corrosion solution due to the production of OH-

which causes an increase of pH value [83]. The corrosion

rate decelerated at pH values above 10.5 because of the

greater tendency for film formation [15].

Furthermore, the microstructure of the alloy caused by

the production process is important. It is reported that grain

refinement decreased the corrosion rate [55, 76, 84, 85].

Equal-channel angular pressing (ECAP) is previously used

to produce homogenous ultrafine grains [76, 84]. The

average grain size of the as-casted Mg–Zn–Zr was

76 ± 5 lm, the as-extruded alloy exhibit a grain size of

2.9 ± 5 lm [55]. In addition, calcium as an alloying ele-

ment also contributes to grain refinement [10]. In a previ-

ous study, a reduction of grain size from 175 ± 15 to

51 ± 5 lm was reported with increasing the Ca content in

Mg–Zn–Mn–Ca-alloy from 0.3 to 1.0 wt% [61].

The velocity of the corrosive media has a considerable

impact on the corrosion rate of magnesium alloys due to

destruction or prevention of a protective film [10, 15].

Waizy et al. [86] developed a corrosion system with a

constant flow rate of Hank’s solution corresponding to the

blood velocity in natural bone. In this in vitro model,

MgCa0.8 screws were tested in synthetic bone via

mechanical pull-out test. The results showed that the pull-

out force decreased of 30 % after 96 h in corrosive med-

ium compared to the non-corrosion group. A maximum

load capacity of 28 ± 7.6 N/h was determined. The bio-

mechanical data suggested that biodegradable screws pro-

vide a promising bone-screw fixation and have great

potential for medical application. A further study by Waizy

et al. [87] was also performed with a constant flow rate. In

this study, orthopedic ZEK100 plates were investigated

J Mater Sci

123

(Fig. 2). After the immersion test, high amounts of O, Mg,

Ca, P, and C were detected on the surface via EDX analysis

which was associated with the precipitation of carbonated

calcium phosphates (Mg,Ca)x(PO4)y(CO3)z(OH)i. The

samples were biomechanically tested via four-point-bend-

ing test [87]. These in vitro studies of orthopedic screws

and plates made of magnesium alloys should be supple-

mented by a forthcoming investigation of intramedullary

nails consisting of LAE442 (Fig. 2).

In conclusion, in vitro studies are useful to investigate

the degradation behavior of new biomedical magnesium

alloys under standardized conditions. The comparison of

the corrosion rate between different alloys is able to pro-

vide valuable information about its corrosion resistance

and behavior. Although in vitro studies attempt to simulate

the biological environment in vivo very closely, the testing

of the degradation in an animal study cannot be omitted

because of the various biological interactions with the

alloy.

In vivo examination of new implants

For the medical or veterinary application of new bioma-

terials, experimental animal testing is indispensable.

Adjacent to the biocompatibility of the newly developed

biodegradable material, degradation rate and mechanical

parameters are of special interest. Different animal models

are available for these questions [88, 89]. The mouse is

commonly used for studies of osteogenesis in bone or soft

tissue [88] and the rat and the guinea pig for simple implant

geometries [8, 90] or fracture models [88]. An alternative

animal model, which combines the easy handling as well as

the possibility to examine orthopedic implant geometries,

is the rabbit [91, 92]. Evaluation of the degradation

behavior of the implant can be performed as well as sur-

rounding tissue reactions [93–95]. In addition, adjacent to

the clinical examination, radiographic and in vivo lCT

examinations can be used to quantitatively assess the

implant degradation as well as changes in bone volume,

bone density, and bone porosity in the direct implant sur-

rounding [96]. After different postoperative observation

periods, histological examinations of the implant interface,

the bone, and the soft tissue around the implant can be

performed. Examination of degradation properties, with

higher resolution lCT and SEM are established methods to

verify the results of the radiographic and in vivo lCT

imaging [96]. For biomechanical analysis of the implant

material, three point bending tests [97–99] and pull-out

tests are available [94]. In addition, examinations of

regional lymph nodes [100] and inner organs [101] can

give evidence of systemic reactions due to the degradation

products and their elimination.

In vivo examination of different magnesium alloys

Whereas specific magnesium-based alloys are already used

in cardiology [102, 103], implant materials for orthopedic

applications in clinical use are not applicable yet. Different

alloys are examined as potential orthopedic implant

material in vivo. As single alloying elements adjacent to

magnesium, calcium [99, 104, 105], and zinc [101] were

tested so far. More than one alloying element is used in

implants with aluminum in combination with zinc [8, 95,

106, 107], calcium [96, 108], or together with lithium and

rare earths [8, 97–99, 109, 110] as in the examined mag-

nesium alloys WE43 [8, 90, 97] and MgZnxMny [101, 111],

ZX50, WZ21 [107], and MgBiCa [112]. Most authors

attest magnesium a good biocompatibility [8, 90, 98, 99,

101, 105, 107] with enhanced bone bonding to implant

surfaces in comparison to conventional used materials [90,

107, 113, 114] and a low inflammatory and immunogenic

potential [90, 93, 100, 101, 107, 111]. So far, only

LACer442 is not recommended for in vivo application due

to a too fast degradation rate in vivo [109]. In comparison

to other degradable materials like polymers, magnesium

alloys exceed in the biomechanical stability [115]. A

negative effect which is described in different studies is the

formation of gas cavities during the degradation process in

vivo. However, no clinical relevance could be observed

Fig. 2 Picture of different orthopedic implant geometries made of magnesium alloys. a MgCa0.8 screw b ZEK100-plate c intramedullary

LAE442-nail

J Mater Sci

123

[96, 116]. Until now, only a single in vivo study exists in

the available literature which describes full degradation of

magnesium-based implants after fixation bone. In this

study, Kraus et al. [117] observed complete bone recovery

even after severe bone affection due to massive gas for-

mation in a juvenile animal model. However, self-healing

properties are less in adults. Therefore, severe bone

affections as described there do not seem to be acceptable

for general orthopedic use. With exception to Kraus et al.

[117] and Castellani et al. [90] who used juvenile rats, all

other studies were performed with adult animal models [96,

99, 101, 104, 118, 119].

Own in vivo research results

In our research group, the magnesium alloys LAE442,

WE43, ZEK100, AX30, MgCa0.8 [96–99, 119], LACer442

[109], AL33 [120], and LANd442 [118] are examined yet.

Therefore, pins of each material (2.5 mm diameter, 25 mm

lengths) were implanted in rabbit tibiae and observation

periods of 3 and 6 months as well as 12 months for

MgCa0.8 and LAE442 were examined with clinical,

radiographical, and histological examinations. In some

alloys (AL33, ZEK100, AX30, and LANd442), in vivo

l-computed tomography with evaluation of changes in

bone volume, density, and porosity in the surrounding of

the implant material was performed (Fig. 3) [96, 118].

Implant degradation was characterized with ex vivo lCT,

volume and/or weight loss after explantation and in some

cases SEM [96, 97, 99, 119].

In addition, regional lymph nodes as well as inner

organs (liver, spleen, and kidney) were examined histo-

logically after implantation of LAE442 and MgCa0.8

implants in comparison to PLA as resorbable and titanium

as permanent implant material [100].

LACer442 and AL33 were excluded for further testing

due to their high degradation rate and their insufficient

biocompatibility [109, 120]. After six months implantation

time, with the exception of ZEK100, all other implant

materials still showed their cylindrical shape. ZEK100 and

LANd442 showed cleft surfaces and losses of implant

material in greater extent than in the other groups [96, 118,

121]. Pitting corrosion was seen as predominant type of

corrosion in MgCa0.8, AX30, LANd 442, and ZEK100

(Fig. 4) [96, 98, 118, 121] in contrast to WE43, which

showed soil-like ablations on the surface and LAE442

which was characterized by homogeneous fissured corro-

sion [97]. All the tested magnesium alloys caused a peri-

osteal increase in the mineral apposition rate, which was

calculated after intravital fluorescent labeling [98, 118]. An

increase in osteoclast activity was seen as well, which was

moderate in the slow degrading alloy LAE442 and higher

in the faster degrading alloys AX30, ZEK100 [98]

(Huhnerschulte, unpublished data). LANd442 caused high

Fig. 3 Trabecular bone formation at the implant interface (redarrow); l-computed tomographic picture

Fig. 4 Pitting corrosion at the surface of an explanted magnesium alloy (stereo-microscopy)

J Mater Sci

123

Table 4 Summary of previous in vitro corrosion tests

Magnesium alloy Corrosive media Immersion time (h) Temperature (�C)

AZ31 NaCl, PBS 2, 24, 48, 72, 96, 144 37 [76]

AZ31 Hank’s solution Up to 480 37 ± 0.5 [84]

AZ31, AZ61, AZ91D, Mg m-SBF 24, 48, 120, 192, 384, 480, 576 36.5 ± 0.5 [130]

AZ31, AZ91, MgCa Hank’s solution, DMEM and

DMEM ? FBS

0, 24, 48, 72, 96, 120, 144, 168 37 [131]

AZ31, AZ91, MgCa Hank’s solution, MEM, MEM ? FBS – 37 ± 1 [70]

AZ31, LAE442 NaCl, PBS, PBS ? albumin – Not mentioned [67]

AZ61, AZ61Ca, AZ91,

AZ91Ca

m-SBF – 36.5± 0.5 [132]

AZ91 Hank’s solution – 20 ± 2 and

36.5 ± 0.5 [82]

AZ91, Mg SBF – 36.5± 0.5 [85]

AZ91 Four solutions with NaCl, NaHCO3,

K2HPO4 and Na2SO4

Up to 168 37 ± 0.5 [78]

AZ91 C-SFB, Hank’s solution, PBS, NaCl,

DMEM

Up to 96 37 ± 0.5 [133]

AZ91 SBF, SBF ? BSA 168 37 ± 1 [74]

AZ91D, LAE442 NaCl, PBS, PBS ? albumin – Not mentioned [67]

AZ91, Mg, WE54 SBF – 36.5± 0.5 [134]

M1A SBF, SBF ? albumin 1-24 37 [73]

Mg SBF 24 37 ± 0.5 [77]

Mg SBF 72, 120, 168, 336, 504 37 [135]

Mg NaCl, NaCl ? HEPES,

NaCl ? NaHCO3, Earle(?), eagle’s

minimal essential medium (E-

MEM), E-MEM?FBS

Up to 336 37 [60]

Mg, Mg–Al, Mg–Ag,

Mg–In, Mg–Mn, Mg–Si,

Mg–Y, Mg–Zn, Mg–Zr

SBF, Hank’s solution Up to 250 37 [127]

Mg, AZ31, MgCa0.8,

Mg-Zn, Mg–Mn,

Mg–1.34Ca–3Zn

EBSS, MEM, MEM ? bovine serum

albumin

168, 336, 504 37 [65]

Mg, AZ91, ZE41,

Mg2Zn0.2Mn

Hank’s solution Up to 312 37 ± 2 [81]

Mg, MgCa SBF – Not mentioned [136]

Mg, Mg–Zn SBF 72, 720 37 [52]

Mg, Mg–Zn SBF Up to 72 37 ± 0.5 [101]

Mg, WE43, E11 Hank’s solution, DMEM,

DMEM ? FBS

Up to 180 37 [75]

MgCa SBF 12, 24, 36, 48, 72, 84, 96 37 [59]

MgCa SBF Up to 250 37 [104]

MgCa Distilled water ? bovine serum

albumin, NaCl, NaCl ? Ab

Up to 120 37 ± 1 [71]

MgCa, Mg–Ca–Y SBF, modified-MEM ? FBS 0, 20, 65, 120 37 [51]

Mg–Mn, Mg–Mn–Zn,

WE43

SBF 24, 48, 96, 216 37 ± 1 [137]

Mg–Mn–Zn Hank’s solution, simulated blood

plasma (SBP)

Up to 288 Not mentioned [13]

Mg–Y–Zn, WE43 SBF (H), SBF (T), MEM, PBS Up to 170 37 ± 2 [138]

Mg–Zn–Ca, Mg–Zn–Mn,

Mg–Zn–Si

Ringer’s solution – 37 [139]

Mg–Zn–Ca–Mn Hank’s solution – 37 ± 1 [61]

J Mater Sci

123

osteoclastic activity although the degradation rate was

moderate in comparison to the other alloys [118, 121].

Owing to profuse bone reactions caused by the implant

material with bone resorption and periosteal bone prolif-

eration, ZEK100, AX30, and LANd442 could not be rec-

ommended for orthopedic application in weight bearing

bones [96, 118]. Promising implant materials after

6 months implantation time were MgCa0.8 and LAE442

[97, 109]. Therefore, longer implantation periods were

performed [99]. After 12 months, MgCa0.8 implants could

only be taken out in sections with exception of one pin with

deep pits of corrosion in contrast to LAE442 which

remained the cylindrical shape to a large extend and

showed fissured corrosion [99]. Bone reactions as sign of

biocompatibility were less in the groups with the slower

and more uniform degrading implant materials, particularly

in the LAE442 group [98, 99].

Systemic reactions could not be observed in any of the

examined degrading magnesium alloys. Even in LANd442,

which is not recommended due to insufficient biocompati-

bility, an increase of IL6 as parameter for systemic inflam-

matory reactions could not be observed [118]. Examined

inner organs (liver, spleen, and kidney) did not show any

histological changes. In regional lymph nodes of animals

with MgCa0.8 and LAE442 implants, non-specific immu-

nological reactions could be found as predominantly foreign

body reactions and were less than in regional lymph nodes of

the clinical accepted biomaterials titanium and PLA [100].

For the use in weight bearing bones, the good biome-

chanical properties are essential as well. Pull-out tests in

the rabbit tibia, with MgCa0.8 as screw material in com-

parison to surgical steel, were performed. MgCa0.8 screws

showed stable pull-out forces during an implantation per-

iod of 4 weeks, followed by a decrease from week six,

which was significant after week eight [94]. It is ques-

tionable if this fast reduction in pull-out forces is sufficient

for the clinical use but MgCa0.8 might be an appropriate

material for screws with less weight bearing. Indeed, the

slower degrading material LAE442 might have a slower

decrease in pull-out forces, and therewith might be an

appropriate implant material even for screws and plates

which has to be examined.

For intramedullary nailing, especially, a high initial

strength is necessary, which should decrease during the

implantation period to avoid stress-shielding effects.

Among the tested materials, LAE442 showed the highest

initial maximum forces in the three point bending test

followed by ZEK100 and WE43 [96, 98].

Summarizing, the alloy LAE442 is the most promising

implant material in our in vivo studies. Following steps

will be to examine orthopedic systems like bolted intra-

medullary nails or screw-plate systems to evaluate the

degradation properties at the interfaces of the material as

well as the bone reactions to higher amounts of degrading

implant material.

It will be the aim of further studies to examine complex

implants made of LAE442 in the sheep as large animal

model. Another approach will be to substitute the rare earth

composition metal by exact defined REE to improve the

reproducibility of the material, and therewith improve the

prediction of degradation rate, mechanical stability, and

biocompatibility.

Acknowledgements The authors gratefully acknowledge the

financial support given by German research society (DFG) within the

collaborative research project (SFB 599). We thank Christopher

Muller for the design of Fig. 2.

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