Cyclic Potentiodynamic Polarisation Curves

Published Date: 02 Nov 2017

1. From cyclic potentiodynamic polarisation curves, self-passivation behaviour with a large passive potential range and low passive current densities for Ti-5Nb-10Zr-5Ta, Ti-10Nb-10Zr-5Ta, Ti-20Nb-10Zr-5Ta alloys was obtained; corrosion potentials exhibited more electropositive values with the increase of the niobium content, due to the favourable influence of this element that, having a nobler corrosion potential acts by its effect of the galvanic couple, ennobling the alloy corrosion potentials.

2. The corrosion and ion release rates have low values that characterise a very good corrosion resistance in the "Very Stable" class; very high values of the polarisation resistances show a very resistant passive film; corrosion rates are lower and polarisation resistances, Rp are higher for the alloy with the highest niobium content, namely, these most protective properties are due to the protective Nb2O5 oxide existing in the alloy passive film and that improves its passive characteristics.

3. Nyquist plots exhibited an incomplete, large semicircle showing a capacitive behaviour, a passive insulating film on the surface of all these three alloys; the semicircle diameters and impedance values increased with the increasing niobium content, indicating a more stable, resistant passive film.

4. Bode phase plots displayed in the low and middle frequency range two phase angles: the first phase angle reveals a typical passive films and a near capacitive response; the second phase angle characterises some relaxation processes at the interface with the electrolyte; these angles have lower values in acid and alkaline Ringer solutions, showing a slightly defective passive film, low dissolution processes through film; comparing the those three alloys, the best angles were registered for Ti-20Nb-10Zr-5Ta alloy, demonstrating a better capacitive behaviour, more protective passive film due to the beneficial influence of the niobium.

5. An electric equivalent circuit with two time constants was modelled: the first time constant is represented by the inner, passive, barrier layer resistance, Rb and capacitance, CPEb; the second time constant is associated with the outer, porous layer resistance, Rp and capacitance, CPEp. All fitting parameters had the most favourable values for Ti-20Nb-10Zr-5Ta alloy as result of its most suitable composition.

6. Open circuit potentials became more electropositive in time and after about 700 immersion hours reached a constant level, indicating the thickening of the passive layer, the increase of its stability. The noblest Eoc values were registered in neutral Ringer solution, proving a very good resistance of the those three alloys in normal functional conditions of an implant. The alloy with the highest Nb content, Ti-20Nb-10Zr-5Ta alloy exhibited the most electropositive Eoc values, namely, the best passivation, confirming the other electrochemical results.

7. Open circuit potential gradients have low values, which can not generate galvanic or local corrosion.

2. Temporary implants Authors: Ecaterina Vasilescu and Paula Drob

In present, titanium alloys are used for hard tissue implant, especially for load-bearing implants as plates, screw, pins, etc. [114-116]. Some metallic implant materials have higher elastic modulus (example: Ti – 105 GPa; Ti-6Al-4V alloy – 115 GPa) [117] than of the bone (10 – 40 GPa) [118] and can cause stress-shielding effects followed by the reduced stimulation of the new bone growth and remodelling, decreasing the implant stability. Another limitation of the current metallic biomaterials is the possible release of the toxic metallic ions or particles through the corrosion and wear process, leading to the inflammations or losses of the surrounding bone tissues. After tissue health, a secondary surgery procedure is necessary to remove the implants. Repeated surgeries increase the cost and create problems for the patient health. The avoiding of the these impediments is the obtaining and using of biodegradable, bioresorbable, temporary implant bioalloys with very good mechanical properties, ossteoinduction and ossteointegration capacity and non-toxicity. Biodegradable metals and alloys have provided mechanical support to bone during its healing and remodelling, while gradually degrading and eliminated from the body. Biodegradable magnesium and magnesium alloys [116,119] are very interesting as temporary implants in the recent years [114,116,120-122]. Magnesium was the first biodegradable implant material used in 1808 by Sir Humphrey Davy [119] because this metal is active and can degrade naturally in the physiological environment by corrosion. Taking into account that for the bone repair and fixture are necessary at least 12 weeks, the temporary implants have slowly biodegrade and maintain their mechanical properties and integrity all this period, and then have slowly eliminate without any discomfort to patient.

2.1. Biodegradable metals and alloys

2.1.1. Biodegradable metals

The most known non-toxic, non-allergic, biodegradable metals are: Mg, Al, Zn, Ca, Li, Zr, rare earth elements (RE) etc.

a) Magnesium

Magnesium [116,119] is an attractive metal for biodegradable implants because of its low thrombogenicity and well-known non-carcinogenity, having important biocompatibility due to the fact that is a structural constituent of the tissues and essential element in the living organism. More than 300 enzymatic reactions are triggered with Mg as part of enzymes or coenzymes and it is essential for the neuromuscular transmittance of the stimuli. It is a substantial intercellular cation which is involved in a great number of the biological reactions of the cells. Magnesium has many similar properties with of the bone [115]: its density is 1.74-2.0 g/cm3 and of the natural bone is 1.8-2.1 g/cm3; elastic modulus is 41-45 GPa and of the bone is 3-30-40 GPa; compressive yield strength is 65-100 MPa and of the bone is 130-180 MPa; fracture toughness is 15-40 MPa.m1/2 and of the bone 3-6 MPa.m1/2. Magnesium is the fourth most abundant cation in human body. The normal level of magnesium in extra cellular fluid is 0.7-1.05 mM/L; when this level is exceeded, many unpleasant phenomena can appear: muscular paralysis, hypotension, respiratory distress and cardiac arrest [122]. By corrosion, magnesium dissolves in human fluid with a high corrosion rate, forming a soluble, non-toxic oxide that is excreted in urine; also, its corrosion leads to the loosing of its mechanical integrity before the bone has sufficiently healed [115] and producing gas as subcutaneous gas bubbles and potential for gas gangrene [120] and reduction of the cell attachment [115]. Also, the increase of the local pH over 7.8 value leads to the alkaline poisoning effect [122,123]. Only ultra pure 99.99% magnesium has a lower corrosion rate [116].

The corrosion reactions of magnesium in the presence of chloride ions from physiological fluid are [114,115]: initially (reaction 1) it forms the magnesium hydroxide Mg(OH)2 that is slightly soluble in water; both magnesium (reaction 2) and Mg(OH)2 reacts with Cl- ions forming magnesium chloride and hydrogen gas (reaction 3).

Mg(s) + 2H2O → Mg(OH)2(s) + H2(g) (1)

Mg(s) + 2Cl-(aq) + 2H2O(aq)→ MgCl2 + Mg(OH)2(s) +H2(g) (2)

Mg(OH)2(s) + 2Cl-(aq) → MgCl2(aq) + H2(g) + 2O2-(aq) (3)

In solution containing chloride ion of 30 mM/L, magnesium develops pitting corrosion; in the human biofluid, the concentration of Cl- ion is 150 mM/L; it results that the magnesium corrodes very rapidly in biofluid with evolution of hydrogen gas.

b) Aluminium

Aluminium in concentration up to 4% can decreases the corrosion rates of the magnesium alloys due to its native protective Al2O3 [116,122] oxide that is more stable than MgO oxide in physiological environment. Some alloys of Mg with Al and Zn enhanced osteoblasts activity in vivo, i.e. no toxicity. However, at higher concentration, aluminium forms Mg17Al12 phase which increases the tendency to pitting corrosion [124]. The concentration limit of Al in serum is 2.1-4.8 μg/L [125]; exceeding this limit, Al can combine with the inorganic phosphates leading to lack of phosphates in the human body, being harmful to neurons and osteoblasts, inducing dementia and Alzheimer’s disease [126,127]. So, it is imperative to control the Al content by the alloying with Mg.

c) Zinc

Zinc [116,122] is an essential element in the human body and improves the strength and castability of the Mg alloys [125], reduces the impurities effects and hydrogen evolution, namely, the corrosion process in its alloys with Mg. The zinc tolerable level in serum is 0.82 mg.kg/day [128]; the overdoses can reduce erythrocyte superoxide dismutase level; normally, the zinc concentration is about 1% in alloys with magnesium.

d) Calcium

Calcium [116,122] belongs to the bone. Low calcium concentration contributes to the solid solution, precipitate strengthening and grain boundary strengthening, acts as a grain refining agent and improves the corrosion and mechanical properties of Mg [115,129].

e) Lithium

Lithium [122] alkalizes the corrosion layer (till pH > 11.5), improving the corrosion resistance and stabilising the Mg(OH)2 film on the alloy surface [130]; but, in vivo alkalinization is a hazard to the human body. Li is used therapeutically in psychosis with a dose of 1300 mg/day. Generally, toxic effect of Li is very unlikely and its effects on the bone are not described yet.

f) Zirconium

Zirconium [122] is the second group of Mg alloys as a grain refiner and it increases the magnesium corrosion resistance. The daily human dose is 125 mg and in higher concentration has toxic effects on liver, lung, breast, nasopharyngeal cancer [131].

g) Rare earth elements

Rare earth elements [116,122] as alloying elements have beneficial influence on the castability, tensile and creep properties and corrosion resistance. As chelated rare earths, these elements are rapidly excreted via urine, but, un-chelated ionic rare earths form colloid in blood, having adverse effects. Some rare earths elements as cerium, praseodymium, yttrium, and lutetium exhibited toxicity to the human body [132].

As was shown above, all biodegradable metals offer some advantages and drawbacks. Thus, it appeared the necessity to improve the behaviour of these metals by alloying with some corrosion resistant metals as Zr and rare earth elements.

2.1.2. Biodegradable magnesium alloys

Many biodegradable magnesium alloys were developed: binary, Mg-Ca, Mg-Zn, Mg-Al, Mg-RE; ternary, Mg-Ca-Zn, Mg-Al-Zn, Mg-Zn-Mn, Mg-Zn-Zr; quaternary, Mg-Al-Zn-Mn, Mg-Li-Al-RE, Mg-Y-RE-Zr; multi-component alloys, Mg-Al-Zn-Mn-Li, etc.

a) Binary magnesium alloys

Mg-Ca alloys with 0.8% or 1.5% Ca content [133,134] showed good compatibility and biomechanical properties for about 23 weeks; in the presence of albumin, the corrosion and hydrogen evolution rates decreased. However, further studies are necessary if its corrosion in time is sufficient for severe internal fracture health, because the implant degradation takes places simultaneously with the bone healing. In higher concentration, from 3.4% to 16.2%, the obtained alloys presented severe localized corrosion [116,121] and a rapid degradation.

Mg-Zn alloys [121,122] having a content of Zn from 1%, 2%, 3%, to 10% exhibited an increased corrosion rate with the increasing Zn concentration; alloy with 3% Zn revealed severe localized corrosion and the alloy with 6.25% Zn also, pitting corrosion; valuable in vitro, in vivo and clinical investigations are necessary.

Mg-Al alloys [116,121] especially with 3% Al concentration developed severe localized corrosion and dangerous effects for human body.

Mg-Re alloys [116,121] containing 5% - 2% Y, 1% - 4% La, 1% - 4% Nd, 1% - 4% Ce showed pitting corrosion and high corrosion rate.

b) Ternary magnesium alloys

Mg-Ca-Zn alloys [116,121] with 0.4% Ca and Zn from 3% to 10% [116,121] registered pitting corrosion and necessitate new studies concerning the quantity of ions released, their limit toxicity doses, the influence of the temperature, the nature and sensibility of the corrosion processes in vivo, etc. Mg-30Zn-4Ca alloy [135] proved higher cell viability and uniform corrosion resistance.

Mg-Al-Zn alloys exist in many variants: Mg-9Al-1Zn [114,119,136,137], Mg-3Al-1Zn, Mg-8Al-0.5Zn [138]; these alloys are characterized by pitting corrosion, self-limited due to the accumulation and stabilization of their corrosion products by OH- ions, limited adhesion and spreading of the cells by the same corrosion products and better corrosion resistance for Mg-8Al-0.5Zn alloy.

Mg-Zn-Mn alloys were studied with the following formulas: Mg-2Zn-0.2Mn [122] that develops a tolerable hydrogen evolution rate, and a good biodegradation process; Mg-1Zn-1.2Mn alloy [139], Mg-0.8Zn-1Mn [140], Mg-1Zn-1Mn [141] without hydrogen evolution and integrity for 1, 4, 9, 18 weeks.

Mg-Zn-Zr alloys were selected by Z. G. Huan et al. [142] with the composition: Mg-3Zn-0.6Zr, Mg-6Zn-0.6Zr, etc; the first alloy Mg-3Zn-0.6Zr showed a low degradation rate, hydrogen evolution, moderate ion release and alkalinization and good cytocompatibility with stimulating effect on bone cells.

c) Quaternary magnesium alloys

Mg-9Al-0.6Zn-0.2Mn alloy [143] had a low corrosion rate for short term and a different one for long term.

Mg-2Al-0.3Mn-0.1Zn alloy [144] reveals a good corrosion resistance in dynamic conditions in 0.01M NaCl solution.

Mg-4Li-4Al-2RE alloys [145,146] did not present gas formation in vivo conditions (rabbit tibia or femora) for 2-12 months.

Mg-4Y-3Re-0.5Zr alloy [146] was implanted in the rabbit tibia and guinea pig femora for 18 weeks and no gas or adverse reactions were observed.

d) Multi-component alloys

Mg-Al-Zn-Mn-Li alloys were studied by Bender et al. [144] concerning the influence of Li and Fe impurities on their behaviour; alloys with low content of Al (1%), the same content of Zn (0.3%) and Mn (0.2%) and different content of Li (11% and 12%), Mg-1Al-0.3Zn-0.2Mn-11Li and Mg-1Al-0.3Zn-0.2Mn-12Li and moderate Li (19%) content, Mg-2Al-0.3Zn-0.2Mn-9Li are not resistant in 0.01M NaCl solution.

Mg-Al-Zn-Mn-X (Ca, Sb, Sb + Bi) alloys [147] were obtained starting from Mg-9Al-0.8Zn-0.2Mn alloy with additions of 0.14% Ca, 0.4% Sb, and 0.4% Sb + Bi; the corrosion rates increased by alloying with above mentioned elements.

Till present, no binary, ternary, quaternary o multi-component alloys satisfy all requirements for a good biodegradable, temporary implant material, especially due to their fast corrosion, degradation rates and hydrogen gas evolution. New research studies are necessary to improve the bio-performances of these alloys.

2.2. Improvement of the temporary implant bio-performances by different protective coatings

The good bio-performances of the Mg alloys, namely, the decrease of their corrosion rates can be obtained by the application of the biocompatible coatings on their surfaces which can act as a barrier against their rapid dissolution. The most favourable coating is hydroxyapatite or its precursors, tricalcium phosphate or calcium phosphate (brushite). Many deposition methods were experimented: electrodeposition, micro-arc oxidation, soaking in different solutions, etc.

2.2.1. Electrodeposition of protective coatings

Electrodeposition of the hydroxyapatite can be realised by anodic potentiostatic electrodeposition [148], or by cathodic galvanostatic electrodeposition [149].

a) Anodic potentiostatic electrodeposition of hydroxyapatite coatings

This method [148] was applied by potentiostatic polarisation at +100 mV (versus saturated calomel electrode-SCE) in a solution containing Ca(NO3)2.4H2O, NH4(PO4)3 and H2O2; an initial brushite CaHPO4.2H2O coating was obtained; then, by the immersion in NaOH solution at 800C for 2 hours, the hydroxyapatite (HA) is formed. This HA coating improved the corrosion resistance of CP Mg in simulated body fluid (SBF).

b) Cathodic galvanostatic electrodeposition of brushite coating

Using mixed solution of Ca(NO3)2.4H2O and NH4(PO4)3 at pH = 5.0, at 400C and a potential of -3.0 V [149] for 4 hours, a coating consisting from flask brushite CaHPO4.2H2O crystallites was obtained; modifying the electrodeposition conditions, a coating with two layers (an inner, dense, fine-grained layer which give better corrosion protection and an outer, well-crystallized, porous layer which provides appropriate pore sizes for bone cell ingrowths) was realised.

c) Micro-arc oxidation coatings

Protective micro-arc oxidation coatings [150] containing non-toxic Mg2B2O5, Mg2SiO4, and SiO2 compounds, were deposited applying a current of 40 mA/cm2, at 250C, for 1-22 min on the Mg surface. Also, on Mg-Ca alloy surface [151], the micro-arc oxidation was applied at different voltages and a coating with good long term corrosion resistance and promoting the cell adhesion, proliferation and differentiation was obtained.

d) Pulse electrodeposition of Ca-deficient hydroxyapatite

By varying the current pulse amplitude and width [152], a Ca-deficient hydroxyapatite (with Ca/P ratio of 1.33 while in HA Ca/p ratio is 1.67) was electrodeposited; this coating showed better adhesion to Mg-Zn-Ca substrate and increased significantly the substrate corrosion resistance, ultimate tensile strength and time of fracture; in vivo experiments are necessary.

2.2.2. Chemical deposition of protective coatings

a) Chemical phosphating coatings

A complex phosphating treatment [153] in a phosphating bath containing Zn(H2PO4)2.2H2O and Zn(H2PO4)2 at 60-650C was applied on the surface of Mg-1.2Mn-1.0Zn alloy and a coating formed mainly from brushite CaHPO4.2H2O and several phosphate compounds Zn2Ca(PO4)2.2H2O and MgZn2(PO4)2 were formed. The brushite layer has been transformed into hydroxyapatite during the immersion in simulated body fluid and released acidic phosphate ions which neutralized the alkalization effect produced by the magnesium corrosion and enhanced the coating biocompatibility.

b) Biomimetic apatite coatings

Hydroxyapatite coatings [154] with two different thicknesses were deposited onto pure Mg substrate by immersion in a concentrated simulated body fluid (SBF) with calcium (3CaPSBF) at 420C for 24 hours; no heat treatment was necessary. The deposited hydroxyapatite coating greatly retarded the corrosion rate and the dual coating was more effective. It results that the degradation rate of Mg can be tailored by the controlling of the coating thickness.

Another biocompatible, biomimetic coating was formed on the Mg surface by incubation in cell culture medium [155]; this coating contained calcium, phosphorus, and carbon ions and reduced magnesium release by 60% compared with non-treated magnesium surface and improved the cell adhesion and spreading.

c) Hydroxyapatite coating with MgF2 interlayer

A double coating [156] with a MgF2 interlayer and HA coating was obtained by fluoride conversion technique and respectively aerosol deposition method. This coating increased corrosion resistance and cell proliferation and differentiation levels on magnesium surface. In vitro and in vivo tests showed that this double coating enhanced the magnesium potential as biodegradable implant material.

d) Aerosol deposition of hydroxyapatite – chitosan composite coatings

The HA – chitosan (a natural polymer derivative of chitin, with excellent biocompatibility, biodegradability, ossteoconductivity and antimicrobial properties) composite coating [157] was deposited on Mg-3Al-1Zn alloy surface by aerosol method. The coating is highly adherent, enhanced corrosion resistance and biocompatibility by the incorporation of chitosan powder into HA powder.

As was presented above, the problem of the biocompatible, biodegradable, ossteoconductive, ossteoinductive, temporary implants is a very actual research theme and many studies and experiments will be necessary to obtain a very satisfactory, temporary implant material.

3. Concept of Ti and Ti alloys as bioactive implants and evolution in understanding benefit and limitations of Kokubo test Authors: Ioana Demetrescu

At the beginning of 1990 J. Biomed. Research have published a paper entitled [158] "Solutions able to reproduce in vivo surface-structure changes in bioactive glass-ceramic A-W3T. The authors were Kokubo, H. Kushitani, S. Sakka, T. Kitsugi, T. Yamamuro. This paper has represented the cause of a large number of investigations and of controversaries as well, but indeed was an important moment in scientifical and clinical progress. The important number of citations of the paper, more than 2000 is due to the fact that its conclusion are based on a very fast and simple chemical experiment for testing biocompatibility. The test is an immersion of bioactive ceramic in various aqueous solutions containing different ions which are supposed to mimic the solution existing in the human body. In the initial research, immersion took place for a period of 7 and 30 days leading to structural changes investigated through FTIR, XRD, SEM, which have been compared with structural changes appeared in vivo.

The authors have concluded that so called buffer Tris solution containing pure water buffered with trishydroxymethyl-aminomethane, and utilised as simulated body fluid is not able to reproduce structural changes in vivo at the surface apatite formation respectively. They have proposed another solution with concentration and pH as in the human plasma, able to reproduce such changes in vivo. The composition of this solution is: Na+ 142.0, K+ 5.0, Mg2+ 1.5, Ca2+ 2.5, Cl− 148.8, HCO3− 4.2 and PO42− 1.0 mM and the solution is buffered with trishydroxymethyl at pH 7.25. An important observation for Kokubo team investigation in 1990 [158,159] is the fact that in vitro experiments need a very careful attention in selection solution able to simulate conditions from human body (SBF). The other papers [160-165] of the Kokubo group have indicated chemical formation of apatite at the surface of vitreous ceramic which involves participation of Ca 2+, HPO42−, and OH− ions existing in body fluids according to reaction:

10Ca2+ +6PO4 3- +2OH- ↔ Ca10 (PO4)6 (OH)2

Apatite formation in this concept is due to catalytic effect of Si-OH groups existing on the surface of ceramic materials immersed in SBF [160] which are able to induce apatite nucleation. Other experiments have introduced the idea that in SBF, groups as Ti-OH, Zr-OH, Ta-OH and Nb-OH could be efficient in the process of apatite nucleation together with functional groups–COOH, -H2PO4-, which are present in SBF [161,164].

All such groups are negatively charged at pH around 7.40. Transmission microscopy (TEM) and X-ray diffraction analysis (XRD) have confirmed that apatite formation took place indirectly via amorphous compounds as calcium phosphates with a low atomic Ca/P.

The fluids from human body are normally considered supersaturated in the apatite precipitated in the presence of bone tissue. This fact is due to the energetic barrier for apatite nucleation which is very high and can be lowered only in the neighbourhood of bone tissue when at the surface are enough efficient functional groups able to nucleate at the surface of the artificial material [165]. Once such nuclei are formed their spontaneously growth is based on consume of calcium and phosphor from bioliquids. Many discussions took place around the equal or not equal efficiency of Si−OH and Ti−OH groups in apatite nucleation and specific arrangements for enhancing efficiency were proposed. The experiments with TiO2 as gel have indicated as more efficient crystalline structure of TiO2 as rutile and anatase which can provide a specific structural arrangement for Ti-OH groups. Further investigations established as superior efficiency in Ti-OH formation the role of anatase [164]

Such observations lead to the development of bioactive materials for bone substitution based on the property of an apatite stratum formation, a stratum which an almost similar structure of bone which is a natural composite with hydroxyapatite and collagen. Having a similar structure such materials present bond ability through bone and initiate osseointegration. Ceramic materials have compared to bone weak resistance to fracture and a higher elastic modulus and this impediment have been solved via bioactivation treatments on materials with high mechanical properties.

The chemical proposed treatment for metallic and ceramic materials with resistance to fracture which involve treatment with NaOH followed by thermal step was applied on Ti and its alloys [162,163,166], on Ta and on ZrO2 treated H3PO4. Such results had introduced the concept of bioactive metals applied in the titanium case [167-169], despite the classic concept of metallic materials as bioinert category.

In this approach of changing materials from bioinert to bioactive via surface treatment, Ti and its alloy which have important resistance to fracture have formed sodium titanate on surface after immersion in 5 M-NaOH solution at 60oC for 24 h, followed by calcination at 600oC for 1 hour [162,163].

In the time of such treatments are taking place simultaneously the following reactions starting with corrosive attack on TiO2 [162]:

TiO2 + OH- → HTiO3 –

Ti + 3OH- → Ti(OH) 3+ + 4e-

Ti(OH)3+ + e- → TiO2 H20 + 0.5H2

Ti(OH)3+ + OH- ↔Ti(OH)4

A later attack on the form hydrated previously produce another hydrated form negatively charged TiO2 * nH2O + OH- ↔ HTiO3– +nH2O which with Na lead to titanate. At a distance of 1 micron from surface, sodium titanate is gradually substituted with metal. It is remarkable that mechanical properties are not affected during this process. Firstly in the environment Ti-OH groups are generated and such groups react with Ca 2+ ions to form amorphous calcium titanate, able to participate in a combination with phosphate ions in order to get amorphous calcium phosphate. This last phosphate could be a precursor of apatite with bone similar structure and Ti could be shortly bonded with living tissue. This osseointegration process was clinical applied starting 2004 for total hip prosthesis [163]. Another important aspect in the behaviour of materials immersed in SBF is connected to the strong negative surface potential [168] at the beginning of immersion in SBF. After a while undergoes a shift to the positive values reaching a maximum, a decrease and a trend to stable values in the negative domain. Each variation of surface charge represents steps of the chemical process as following:

(1) formation of Ti-OH groups negatively charged during the change of Na ion from titanate with H3O+ from the fluid ;

(2) formation of amorphous calcium titanate stratum positively charged after combination of Ti-OH (with negative charge) with Ca 2+;

(3) formation of amorphous calcium phosphate after combination of calcium titanate positively charged with phosphate ion;

(4) formation of negatively charged crystalline apatite.

According to the above steps apatite mechanism formation is based on electrostatic interaction between metals and the ions existing in biofluid [168]. Publication of spontaneous „Formation of Bonelike Apatite Layer on Chemically Treated Titanium Metals [170], have extended the procedure of metal coating with phosphate using SBF solution, despite the fact that proposed initial SBF solution has not the real composition of human plasma and various SBF with improvements have been used by different researchers groups during the time [171-174]. A revised simulated body fluid [173] was proposed in 2003 by Kokubo group taking into account that a part of calcium and magnesium can not participate at the process of apatite precipitation due to the fact that are bonded to protein.

In 2006 a red paper in Biomaterials [174] journal introduced a very simple and convenient test in evaluating bond ability to bone of any material taking into consideration that the main property of a material in order to function as a bone substitute is such ability. The test does not included animal experiments which at the level of 2006 were already restricted and involved special expensive facilities. Such principle was supposed to be applied on any material leading to a clear result after a short period of time. It is to point out that metallic materials from titanium family were partially more extended due to this test.

The test was confirmed on a large variety of materials exploited efficiently in the conditions of bond ability to bone and at the end of 2006 SBF Kokubo became a general procedure to evaluate material bioactivity based on capacity to induce formation of an apatitic stratum in vitro. Kokubo test was registered as standard [175] entitled „Implants for surgery — In vitro evaluation ISO 2007 for apatite-forming ability of implant Materials, with number ISO 23317 standard (Kokubo & Takadama 2006). The standard has affirmed that is a need for animal test for a final evaluation. A lot of cell culture experiments with positive result on ceramic materials [176-178] as Al2O3, ZrO2, TiO2, and Mg2SiO4 considered bioinert due to the fact that are not able to induce apatite formation lead to the many controversaries regarding the possibility to use the test as a general method [179,180]. Adhesion, proliferation and cellular differentiation considered as a useful tool in bioactivity evaluation become more important and the number of papers with criticism regarding Kokubo test [179-181] increased leading to conclusions as following: the procedure is not concludent taking into account only apatite formation from supersaturated solutions without an analysis of the biologic process [181]. According to the authors [181] bioactivity prediction should be based on osteoblasts behaviour in vivo or vitro [181]. An indication about how to use the test carefully is a result of controversaries, as well as the last paper of Kokubo group about his test limitations [182]. In this paper is described an experiment trying to explain the main factor governing the capacity of apatite forming on Ti. In this experiment the metal was kept in HCl or NaOH solution having different pH varying from 0 to 14 and after that was thermally treated at 600°C.

The apatite formation have been observed on the surface of Ti only in the conditions of immersion in SBF at pH lower than 1.1 or higher than 13.6; for experiments performed in solution with intermediate pH values the apatite formation have not been observed, probably due to the fact that in such conditions the surface charge is neutral. The ability to apatite formation being related to positively or negatively surface charge respectively, such conditions could be reached only in strong acidic or basic solution. As a conclusion we can say that there are limitations in using Kokubo standard, depending on pH solution.

Any way, after immersion in SBF solution a protective layer appears and a better stability and less ions release are expected. Despite the limitations the test is an useful procedure in improving metallic implant behaviour as can be seen in some recent papers [183,184].

4. Antibacterial effect of various implant coatings at micro and nanolevel Authors: Cristian Pirvu and Camelia Ungureanu

Implant surfaces in internal fixation are generally designed to favour the adherence of soft and hard tissue and to lead ultimately to tissue integration and osseointegration. In some areas minimal bone bonding is desired especially for implants that require removal. Soft tissue infection and ossteomyelitis are serious complications associated with implants .

Bacterial infection of implanted medical devices is otherwise a serious problem in the biomedical field. Postoperative infection associated with implants remains also a serious complication in orthopaedic surgery . Thus, inhibiting bacterial adhesion is a crucial step to preventing the implant-associated infection. Implant-associated infections are the results of bacteria adhesion and the biofilm formation .

Biomaterials like Ti-based alloy, Co and its alloys are used as bone plates and dental implants due to their excellent mechanical properties and endurance but long-term performance of surgical implants is also directly depending on their antibacterial effect [188].

Several biomaterial surface treatments have been proposed as a means of reducing the incidence of implant-associated infections.

In fact biomedical coatings used to improve the tissue-biomaterial interface should be viewed more complex aiming at more features of these biomaterials such as bioactivity and biocompatibility, anticorrosion and anti-wear properties and antibacterial activity (Fig. 12).

Fig. 12. Requirements for various implant coatings

The problems in the osseous healing of implants appear to be resolved but the adsorption of biomolecular pellicles and the subsequent accumulation of bacteria on these surfaces are still the main factor for the induction of inflammatory processes. Thus, although the biocompatibility of Ti has been confirmed in many papers , it is still difficult to meet all the requirements of the versatile biosurface, such as antibacterial ability, biocompatibility, osseointegration, and mechanical properties .

Considering that the adhesion and the colonization of bacteria to an implant surface are considered to play a key role in the pathogenesis of infections, several researchers have been attempting to develop different antibacterial coatings which should correspond to increasingly diverse requirements, as can be seen in the figure 12.

a) Antimicrobial polymers coatings

Antimicrobial polymers as polymeric biocides, is a class of polymers with antimicrobial activity, or the ability to inhibit the growth of microorganisms such as bacteria, fungi or protozoans. The research involving studies on antimicrobial polymers has significantly increased in the last ten years ; antibacterial polymers can be added to surfaces without losing their biological activity and become important as an alternative to antibiotics.

The studied literature reveals three general types of antimicrobial polymers: polymeric biocides, biocide-releasing polymers and biocidal polymers. Further biocide release systems in general and microbe repelling surfaces are discussed in others reviews .

The first models were proposed in 2001 by Tiller and show how surface attached antimicrobial polymers act as a contact-active surface .

Page et al. propose a concept of the polymeric spacer effect which presumes that a surface grafted biocidal polymer might be capable of penetrating the bacterial cell wall of an adhered bacterial cell . If it reaches the cytoplasmic membrane, the cell it can be killed by disruption of the phospholipid bilayer. Another mechanism proposed that the hydrophobic surface and the negative surface net charge of microorganism lead to an attraction of the cells to cationic surfaces and destroys their cell wall.

In the recent years using of conducting polymers, termed ‘fourth-generation polymeric materials’, has become increasingly in literature both as polymer coating or polymer composite .

As a result of these investigations, polymers and their composites/blends were found to be useful as a new type of antibacterial agent, self-clean as well as multifunctional material for improving the human health and living environment.

b) Silver antimicrobials coatings

Metal nanoparticles with antibacterial activity when embedded and coated on to surfaces can find immense applications in biomedical and surgical devices. Silver compounds and silver ions are known to show antimicrobial properties and have been used in a wide range of applications. Such coatings could be more important taking into account that antibacterial treatments with antibiotics are becoming less effective due to their use. Ag+ ions have a significant antibacterial effect and have found uses in a significant number of applications .

The treatment with the silver ions results in similar morphological changes in both the Gram positive and Gram negative bacteria. The inhibitory activity of silver ions is higher in case of Gram negative bacteria . It is widely believed that the effectiveness of Ag+ as an antibacterial is due to its ability to bind with thiol (–SH) groups on proteins and enzymes . One possible disadvantage of silver based antimicrobials is the possibility of Ag ion cytotoxicity towards mammalian cells, as Atiyeh et al. reported .

Demetrescu et al. talk about merit and demerit aspects in the performance of the other coating based nAg - hydroxyapatite deposited on TiAlZr alloy. The merit aspect is related to the increase of the stability in biofluids and to the enhancement of the antibacterial effect. The demerit part consists of inducing the decrease in viability and osteogenic activity of the cell line due to the presence of Ag nanoparticles .

Silver ions prevent DNA replication and affect the structure and permeability of the cell membrane and have also been shown to interact with DNA to enhance pyrimidine dimerization by the photodynamic reaction and possibly prevent DNA replication. Sharma et al. has been demonstrated that silver nanoparticles retain their bactericidal properties when included in different coatings. The amount of silver content in the coatings is a very important factor that affects the coating antibacterial properties. A high silver concentration leads to poor biocompatibility while a low silver concentration results in poor antibacterial properties. An appropriate silver content should be chosen to balance the biocompatibility and antibacterial properties of the coatings.

The size of the particle plays an important role in antibacterial activity . The antimicrobial activity was particle size dependent. Small particles exhibited higher antimicrobial activity than big particles. The antibacterial properties are related to the total surface area of the nanoparticles. Smaller particles with larger surface to volume ratios have greater antibacterial activity .

c) Copper antimicrobials coatings

Since ancient times, the copper ion has been well known to be effective against a wide range of bacteria and virus. Nowadays, different studies use many products with copper ion made for having antibacterial effect such as medical applications. Thus, Weaver and Finke have estimated the antimicrobial activities of copper versus stainless steel and have reported a number of microorganisms of clinical importance, including MRSA and E. coli. In all experiments, stainless steel was shown to be inert compared to copper; the copper surfaces clearly shown an antimicrobial effect.

At the same time CuO may find potential application as antimicrobial agents . CuO nanoparticles were effective in killing a range of bacterial pathogens involved in hospital-acquired infections ; Copper ions released may also interact with DNA molecules and intercalate with nucleic acid strands. Copper ions inside bacterial cells also disrupt biochemical processes .

Copper exerts its toxicity by generating reactive oxygen species through copper meditated redox cycling. Although, the chemistry and biochemistry of copper were simply and clearly reviewed by Parker (1981) and by Hughes (1987) , the exact mechanism behind bactericidal effect of copper nanoparticles is not clear.

d) Titanium dioxide antimicrobial coatings

The bacteriostatic agents such as titanium dioxide (TiO2) nanoparticles have also attracted a great deal of attention in the recent time. Titanium dioxide is widely used as a self-disinfecting and self-cleaning material in many applications. Thus, photoactive pigments such as TiO2 have been used for antimicrobial purposes. Generally, TiO2 reacts with light and creates hydroxyl radicals from organic particles and cause an oxidative or reductive effect that degrades microorganisms . The efficacy of TiO2 semiconductor particles as a means of antimicrobial agents was first realized by Matsunaga et al. in 1985 . In this first study it was found that platinised TiO2, when irradiated with ultra band gap UV radiation acted as an antimicrobial agent being 100 % effective against E. coli after 2 hours of UV illumination. Illumination of TiO2 generates excess electrons in the conduction band (e-) and positive "holes" in the valence band (h+). At the TiO2 particle surface the holes react with H2O or surface OH- groups to form HO- radicals. Excess electrons in the conduction band react with molecular oxygen to form superoxide ions .

The previous scheme describes the production of reactive species on the catalyst surface. The hydroxyl radicals produced by the redox process at the TiO2 surface are highly reactive and non-selective. These attributes make the radical species an extremely potent biocide

The three competing theories were evaluated and considered in the light of the latest evidence collected: 1. direct oxidation of coenzyme A, which inhibits cell respiration; this was the original theory proposed by Matsunaga et al.; 2. cell wall decomposition observed by transmission electron microscopy; 3. cell wall damage followed by cytoplasmic membrane damage.

Lu et al. examined the effect of TiO2 on E. coli, using AFM and measurement of K+ ion leakage. AFM was able to show the decomposition of the cell wall, followed by the destruction of the cell membrane. Cell death was due to leakage of the cytoplasm through the damaged membrane. K+ ions are vital to bacterial cells as they play an important part in protein synthesis. AFM show that the cell membrane is compromised by the action of TiO2 under photocatalytic conditions.

Matsunaga et al. compared the antimicrobial effect of the UV light by itself, or the synergistic action of the UV light and the photocatalyst and no microbicidal action were observed in the absence of the photocatalyst. In recent years, visible light absorbing photocatalysts with Ag/AgBr/TiO2 has proved to be successful at killing S. aureus and E. coli .

Nanostructured titanium oxide is a heterogeneous catalyst whose efficient photoinduced activity, related to some of its allotropic forms, opened different ways for its widespread technological use .

e) Antibiotic coatings

Organic antibacterial agents include antibiotics, such as tobramycin, gentamicin etc or different plant extracts.

The antimicrobial activities of plant extracts may also reside in a variety of different components, including aldehyde and phenolic compounds. Naturally occurring combinations of these compounds can be synergistic and often results in crude extracts having greater antimicrobial activity than the purified individual constituents.

Stigter et al. used a biomimetic co-precipitation approach to incorporate an antibiotic (i.e., tobramycin) into hydroxyapatite coatings aiming at preventing post-surgical infections in orthopaedic. Tobramycin was chosen because of its broad spectrum against most Gram-negative bacteria and some Gram-positive bacteria and can prevent growth of E. coli, a Gram-negative bacillus frequently responsible for post-surgical infections in orthopaedic surgery.

The results demonstrated the efficacy of the biomimetic coatings combined with tobramycin to prevent local post-surgical infections .

f) Diamond-like carbon (DLC) coatings

Liu et al. reported reduction of bacterial adhesion on modified DLC coatings (DLC is described as a biocompatible surface coating for biomedical devices such as stents or replacement joints . DLC film can reduce the adhesion of various microorganisms to a stainless steel substrate, and by doping the DLC with Si or N, this can be reduced even further [235]. This represents a very interesting and new area of research, and doped DLC coatings may be very useful to prevent infections after various implant.

g) Bacteriophage coatings - future directions in implant surface modifications

Bacteriophages are viruses that infect prokaryotic cells. Phages usually target individual species of bacteria, bind to their surface, inject the genetic material and replicate within the bacterial host and the result is the lysis of the host cell, and the release of more phages .

The concept of modifying a surface with bacteriophages in order to produce an antimicrobial surface coating is a very recent development . Curtin et al. demonstrated successfully on a hydrogel - coated silicone catheter model in 2006 .

Curtin et al. monitored the formation of a Staphylococcus epidermidis biofilm on catheters pre-treated with coagulase-negative staphylococcus phage 456. It was shown that this modified surface, containing phage units, reduced biofilm formation significantly.

There are, however, a number of potential drawbacks. The most important is the inherent specificity of the phage for each bacterial species. It is less useful for a surface, where a number of different microorganisms, may be present. Phage-treated surfaces would have to be monitored and their formulation modified to remain efficacious-which will no doubt cause problems for regulatory approval.

Therefore it can be concluded that different types of coatings can be applied to improve the quality of the implants and protect them against bacterial infection and the merit and demerit effect must be as well evaluated for optimal choice.

5. Nanodimension effect in bioperformance of implant biomaterials Authors: Ioana Demetrescu

Despite the fact that usually in biomaterials science „bioactive" represents a merit size, there are situations where bioactivity is a demerit [240]. Bioactive surfaces presented bone bonding ability, but some of processing procedures in order to get bioactivity have the potential to form residuals which can be harmful to the cell. Such approach was completed with the introduction of the term „bioreactive" for both merit and demerit aspects of micro/nanomaterials behaviour. However, recently the elaboration of controllable nanoscale structure at the bioalloy surface seems to be an alternative in improving cell adherence and proliferation, due to the fact that the nanoarchitectural features of the cell substrate determine cell shape and contacts [241].Therefore, they are important for cell growth, gene expression and cell differentiation and as examples our studies have found that cell behaviour on Ti implant alloy is critically dependent on different TiO2 nanodimensions as the diameter of nanotubes grown elaborated at surface [241,242]. In their studies, Park et al. [243] demonstrated that the maximal cell response in terms of cell adhesion, proliferation, migration, and differentiation occurs at a nanospacing of 15 nm and decreases on 70- and 100-nm nanotubes. This cell behaviour was explained by a correspondence of this nanotube spacing with the lateral spacing of integrin receptors (about 10 nm) in focal contacts on the extracellular matrix.

Regarding the unintentionally toxicology effect as a size effect of nanomaterials [244], this one was intensively discussed after the rapid development of nanotechnology, when nanolevel became interesting for tissue regeneration and a large variety of nanostructures with well-defined porous microchannels as biocompatible supports for culture growth was described in the literature [245-251].

Despite the fact that nanosizing effects of materials on biological organisms was intensively investigated by biochemical cell, functional tests, cell proliferation and animal implantation testing, this effect is not totally known and needs more experiments and discussion looking for both positive and negative influence on behaviour at nanoscale. It is to point out that the problem of size dependent properties of nanobiomaterials was identified, but the critical size for each kind of nanoarchitecture is far away from the actual knowledge needs in the field, and could be a step in reaching more bioperformance.

There are three kinds of promising nanoarchitectures in our century, the nanowires, nanotubes, and nanoparticles. The mostly investigated nanotubes are carbon (both single and multiwalls and TiO2) and the most studied nanoparticles are silver and TiO2, Au and Fe. The approach at 1D, 2D and 3D level for a critical size with best performance is original and innovative, justifies the interdisciplinary side and relies on the synergy of ascertainment of conceptual methods using the specific means for each fundamental science, respectively chemistry, physics and biology. This subject is a relevant one in the context of the new relationship between biomaterials and nanotechnology [252] which enlarged the definition of the biomaterials [253] and leads to fundamental researches related to the ability of a nanostructure with the same composition, to increase or to decrease a performance as a function of a specific dimension.

Biointeractive term was used for nanomaterial in the idea that the size itself induces reaction to cell and tissue. Generally speaking, the nanosizing physical effect consisting in increasing of specific surface area will induce more chemical activity and reactivity and that means an enhancement of cell adherence and proliferation, but in the same time it means more ion release with toxicity acceleration. Conversion of function by nanosizing may appear as well [254]. In other words, both merit and demerit aspects may appear as aspects of bilateral nature of biocompatibility [1]. Nanosizing may promote conversion of functions and a clear example is Ti behaviour from its biocompatible nature in macroscale to inflammation, leading to osteolysis, in abraded micro/nanoparticles produced from an artificial joint [255,256]. These phenomena cannot be explained by the specific surface area effect and understood to be a different effect, i.e. physical size and shape effect, apart from the material properties of either. As an example the TiO2 nanotubes field on implant alloys has greatly evolved [241]. So far, there have been four generations of titanium oxide nanotubes obtained via anodisation. In the first generation, TiO2 nanotubes were elaborated in HF-based aqueous electrolytes, or with some other inorganic acids and fluoride salts, such as H3PO4/NaF, NH4F/(NH4)2SO4, and Na2SO4/ NaF [257]. Monitoring the pH of the electrolyte, the second generation of nanotubes was fabricated and tube length was increased to a few micrometers [257]. In the third generation, using a mixture of organic and inorganic or an almost water-free polar solution [257], tens even hundreds of micron-length nanotubes were prepared. Effects of a mixture of electrolytes (inorganic + organic) is related to the lower chemical dissolution in comparison to aqueous electrolytes, the formation mechanism of nanotubes being different, especially due to the change of viscosity [257]. The last generation of anodic titanium oxides nanotubes are fabrication via two or more steps and leads to nanotubes multilayers [258]. As a result of all the extensive research regarding TiO2 nanotubes, it is possible to tailor the dimensions and morphology of the nanotubes changing their bioperformance as a result of nanodimension effect [259,260].

Despite the good behaviour of an implant in simulated conditions a part of bioperformance, is a need to discuss antibacterial effect as a function of nanoparticle dimension introduced in system for their antibacterial action. .Bacterial infections after introduction of an implant to the body are usually caused by the adherence and colonization of bacteria on the surfaces of implants [261]. Over the last few years, new and old strategies have been proposed to control and prevent microbial contamination of implants. A part of such strategies involve silver use as an antibacterial agent [203]. Silver had beneficial healing and anti-disease properties and gained approval as an antimicrobial agent prior to the antibiotics period. With the introduction of antibiotics in the 1940s, the use of silver as an antimicrobial agent diminished, but before the end of the century has been a resurgence of the promotion of silver as an effective antimicrobial agent despite controversaries on the subject. The last decade introduced the same trend and silver nanoparticles were intensively studied for the effectiveness in preventing infection [262]. The silver particles diameter seems to be key factor in understanding their antibacterial activity and in this idea based on our expertise in the field our last papers have promoted the idea that smaller particles has smaller antibacterial effect having less space available for bacteria colonization [263]. We are trying now as future perspectives to find critical size for best performance and to propose a model for such biointeraction with specific bacteria. Understanding the behaviour at biointerface as a result of a critical nanobiomaterials size permits to propose a model of such biointeractions and is a motivation for future achievements in bioscience and tissue engineering.