Modificarea suprafetelor implantabile, in vederea cresterii bioperformantelor.

Modificarea suprafetelor metalice implantabile cu acoperiri biomimetice, puncte forte si puncte slabe.

Partea 2.

I. DemetrescuCurs POSDRU

University Politehnica of Bucharest, University Politehnica of Bucharest, ROMANIAROMANIAFaculty of Applied Chemistry and Materials ScienceFaculty of Applied Chemistry and Materials Science

Motivatie si suport

• Motivatie : necesitatea de a imbunatati biomaterialele implantabile in contextul dezvoltarii acoperirilor de suprafata

• Suport : Proiecte si colaborari:• C.N.C.S.I.S.TIPA. „Obtinerea si caracterizarea de noi micro si nanostructuri compozite cu utilizare in

ingineria tisulara”• Elaborarea si testarea in vitro si in vivo a unor elemente de protezare pentru ortopedie, realizate din

noi biomateriale romanesti • Bilaterala Franta Brincusi «Couches minces d'oxyde d'aluminium et d'oxyde de titane pour

différentes applications technologiques et biomédicales• Proiect CEEX Micro si nanostructuri obtinute prin bioactivare chimica si electrochimica cu aplicatii in

medicina regenerativa• Proiect PN2 IDEI Studii exploratorii asupra mecanismului de formare si inducere de noi proprietati

unor electrozi modificati cu forme structurale TiO2 nanotuburi / nanoparticule si compozite polimerice

• Proiect PN2 IDEI complexe PCCE Noi concepte si strategii pentru dezvoltarea cunoasterii unor

noi structuri biocompatibile in bioinginerie

Synthetic bone-like materials with composites of natural and polymeric materials with HA

•Recently were fabricated composite nanofibrous substrate of Chit/HA (chitosan/HA - 80:25) prepared by dissolving in TFA/DCM (trifluoroacetic acid/dichloromethane) (70:30, w/w) for 5 days and electrospun to fabricate a scaffold for bone tissue engineering. HA (25 wt %) was sonicated for 30 min to obtain a homogenous dispersion of nanoparticles within the Chit (80 wt %) matrix for fabricating composite nanofibrous scaffold (Chit/HA).

•The nanofibres of Chit and Chit/HA were obtained with fibre diameters of 274 ± 75 and 510 ±198 nm, respectively, and characterized by FESEM (field emission scanning electron microscopy) and FTIR.

•The interaction of hFOBs and nanofibrous substrates were analysed for cell morphology (FESEM), mineralization [ (Alizarin Red-S) staining], quantification of minerals and finally identified the elements present in Chit/HA/osteoblasts by EDX analysis, which confirmed that the spherulites contain Ca and P the major constituents in calcium phosphate apatite, the mineral phase of the bone. Mineralization was increased significantly up to 108% in Chit/HA compared with Chit nanofibres. The electrospun Chit/HA nanofibrous substrate is a potential biocomposite material for the mineralization of hFOBs. required for enhanced bone regeneration. Cell Biol Int. 2011 Jan;35(1):73-80

containing the cell attachment sequence arginine–glycine–(aspartic acid) (RGD) are fabricated through simple deposition of the ELP dissolved in aqueous-based solutions. The biopolymer is produced and characterized using electrophoresis and mass spectroscopy. The temperature and pH responsiveness are assessed by aggregate size measurements and DSC. The deposition of the studied ELP onto chitosan is followed in situ with a quartz-crystal microbalance with dissipation monitoring (QCM-D). Contact angle measurements are performed at room temperature and at 50 OC, showing reversible changes from a moderate hydrophobic behavior to an extremely wettable surface.

Smart thin coatings using a recombinant elastin-like polymer(ELP)

Water drop profiles on ELP-coated substrates at 25 and 50 oC.

b) Contact angles on glass (&), chitosan monolayer (*), H-RGD6 monolayer (D), and chitosan/H-RGD6 coatings (^). Error bars represent one standard deviation.

c) Contact-angle profile as a function oftemperature for H-RGD6 coatings.

Advanced Functional Materials 2009

H-RGD6 is a recombinant ELP that possesses charged residues,providing weak acidic or alkaline characteristics to the biopolymer

•The ability for H-RGD6 to deposit onto chitosan was first monitored by QCM-D, a technique able to detect masschanges of the order of ng /cm2 and measures the viscoelastic properties of the resulting surface.

•The frequency variation, Df, decreases with time, due to the deposition of polymer mass on the surface of the crystal.

•On the other hand, the dissipation, DD, increases with time,revealing that the film is not rigid and begins dissipating energy, thus exhibiting the typical viscoelastic behavior

•QCM-D monitoring of frequency (Df) and (DD)dissipation changes obtained at three different harmonics:5th, 7th, and 9th, during a)

deposition of chitosan (step 1) and H-RGD6 (step 3), and c) deposition of H-RGD6 (step 3) on a bare crystal (step 2 relates to rinsing). Plotsrepresent: Df5 (D), Df7 (&), Df9 (*), DD5 (~), DD7 (&), DD9 ().

(b) and (d) show the estimated thicknessevolution with time ofH-RGD6 onto chitosan and bare

crystal, respectively,in which time zero

corresponds to the initial moment of the passage of the H-RGD6 solution over the substrates..

AFM image (tapping mode) of H-RGD6 coated surfaces at a) 25oC and b) 50o C..

•AFM analysisperformed at room temperature reveals a smooth surface and no organizedstructure. •At 50 o C, the surface presents spherical nanometer-sized structuresof collapsed biopolymer chains.

•Such results suggest that the ELP chains,when collapsed, aggregate into micelle-like structures at the surface of the substrate, increasing its water affinity.

• The ability of H-RGD6 and RGD(-) topromote cell adhesion in comparison with chitosan surfaces was assessed at different time points. Cell culture was performed at 37 oC, at which the surface is moderately hydrophobic. Cell proliferation and activity was assessed through DNA quantification .

• After 24 h the DNA content was significantly different between these substrates. There are significant differences in the surfaces coated with this biopolymer for 4 days and 7 days, in comparison with chitosan or RGD(-), showing that H-RGD6 coating also increased cell proliferation during the studied timeline. DNA quantification in cell lysates on TCPS

(Tissue culture polystyrene) chitosan, H-RGD6, and RGD(-), for culture periods of 24 h, 4 days and 7 days.

• Cell proliferation and activityassessed through ALP assay

• Secretion of ALP is an important indicator determining the activity of the cells.

• The statistical analysis of the data shows differences in Saos-2 adhesion to H-RGD6 and RGD(-) versus chitosan at 24 h.

• After 7 days the differences in cell viability become more evident in the surface with H-RGD6 in comparison to RGD(-) or chitosan. ALP assays data on chitosan, RGD(-), H-RGD6

and TCPS.

Differences in the number of cells attached to the different surfaces were also observed by DAPI staining(4',6-diamidino-2-phenylindole is a fluorescent stain) , with more SaOs-2 cells(sarcoma osteogenic ) observed in surfaces coated with H-RGD6 and RGD(-), such kind of smart coatings may be used to control cell adhesion as a response to temperature.

• QCM-D results showed that stable, robust coating of about 30nm could be generated over chitosan simply based on the synthetic protein adsorption.

• The straightforward method should, in principle, enable us to coat substrates with complex geometry,including scaffolds or particles exhibitstrong variations in their surface properties, (wettability, topography), depending on T and pH.

• At room T, the surface was more hydrophobic and smooth.

• For warmer conditions, the surface is very hydrophilic with organized spherical structures due to the biopolymer-chain collapse.

• We hypothesized that the biopolymer exposed the charged residues towards the outside and the hydrophobic chains towards the inside,

and that such conformation increased the hydrophilicity of the ELP-coated surfaces,

DAPI staining of the substrates: chitosan (a, d, g), RGD(-) (b, e, h), and H-RGD6 (c, f, i). Tests at 24 h (a, b, c), 4 days (d, e, f ), and 7 days (g, h, i). Scale bar: 50mm.

Human Elastin-derived Biomimetic CoatingSurface to Support Cell Growth (International Journal of Biological and Life Sciences 8:3 2012)

• Mimicking the structure of naturally occurring proteins by using artificial protein polymers with specific materials properties, a new sythetic gene coding for a Human Elastin-Like Polypeptide was constructed and expressed.

• The recombinant product was tested as coating agent to realize a surface suitable for cell growth.

• Coatings showed peculiar features and different human cell lines were seeded and cultured.

• All cell lines tested showed to adhere and proliferate on this substrate that has been shown also to exert a specific effect on cells, depending on cell type.

• The first macromolecule, named HELP (Human Elastin-Like Polypeptide) set up was based on part of the sequence of exon 23, coding for a crosslinking domain rich in alanine and lysine and on exon 24, based on the hexapeptidic VAPGVG repeat HELP1, was realized assembling only the sequence coding by exon 24, giving an expression product lacking the alanine/lysine-rich crosslinking domains

HELP1 coated surface. A, phase contrast microscopy; B,scanning electron microscopy, bar is 50 μm

Phase contrast microscopy of Ea.Hy926 (A and B), MCF-7 (C and D) and A549 (E and F) cells cultured on control standard treated plastic (A, C, and E) and on HELP1 coating (B, D, and F) at 72 hours after seeding.

The viability and proliferation assays performed on the cultures on HELP1 coated surface

•

Proliferation assay of the different cell lines cultured onHELP1 coated surfaces as percent of the control culture on standard treated plastic. Ea.Hy926 (white bar), MCF-

7 (black bar) and A549 (grey bar) at different times.

• Ea.Hy926 cells (endothelian ) gave a signal that resulted even higher than that of the control counterpart. This suggests that HELP1 coating may have apositive effect on Ea.Hy926 cell proliferation. The cultures after 72 hours show a tendency to reach a plateau value lower than that of the controls.

• This assay, being related to metabolism, did not distinguish between viability and proliferation, the data could reflect the fact that responsive cultures did not occupy all the surface, but have preferential areas of proliferation; cells became more crowded, reaching confluence very quickly and stop dividing.• All the cultures showed good vitality after one week inculture,showing that the new recombinant HELP1 is suitable as coating.

• Cells of different origin are able to adhere and grow with negligible cytotoxicity; depending on cell type, HELP1coating has been shown to elicit a cell response. •biopolymers from synthetic genes could be applied in development of smart biomimetic• surfaces for cell growth.

Sample Ecor (mV/SCE)

Ipas (µA/cm2)

Icor (µA/cm2)

Vcor (mm/Year)

Ti -131 26 4.81.10-6

4.17.10-4

Ti-col 617 25 3.17.10-6

2.75.10-4

J.Mat. Science Mat. Medcine,18,10 2075- 2083,

2007)

Control Ti Ti-col

Viability 100 74.3 70.0

Cell spreading

13.9 8.5 11.7

Ti treated with collagen

Strong points in phosphates deposition

• The coated material is more ductil and the surface is bioactive.

• The faster process of adap implant to tissue in the neighberhood

• a reduction in healing time • enhancing osseointegration process • a reduction in the amount of ion release .

Methods of phoshates coatingDeposition method Strong points Weak points

Thermal Spraying (Pulverizare termica)

High deposition rate High Temperatures, risck of raw material decomposition,At fast cooling the result are amorphous film

Sol - gel Can be coverd complexe structureslow processing temperatures low Low cost

High Temperatures for sinteringProblems due to temperatures strss

Pulsed Laser Deposition(cu impulsuri laser)

The result is a film with an uniform thickness high costs Long elaboration time Can not be coverd complexe structuresthe result is amorphous film

Dip Coating (Prin imersare) Low cost

Short coating time Can be coverd complexe structures

High Temperatures for sinteringProblems due to temperatures stress

Electrophoretic Deposition The result is a film with an uniform thickness High deposition rate

Can be coverd complexe structures

craks appear frequently is a need for High Temperatures for sintering

Sputter Coating (prin descarcare de electroni)

The result is a film with an uniform thickness high costs Long elaboration time Can not be coverd complexe structuresthe result is amorphous an film

Electrochemical deposition

Strong points

• Low temperatures for processing • Uniform deposition on irregular surfaces • Low cost ;• Good ratio quality/price ;• Easy to monitorize microstructure coating, changing

working parameters ;

Parameters which play role in changing depositionrate and morphology

• Electrolyte concentation• Electrolyte pH;• Voltage ;• Ionic strenghts of electrolyte and of ionic species

presented( CaCl2, Ca(NO3)2, Ca(OCOCH3) + NH4H2PO4; Ca(H2PO4));

• Electrolyte temperature;• Agitation solution level ;

Electrochemical parameters of Co-Cr-Mo alloy untreated (UN) and treated with collagen gel (TC) in artificial saliva Duffo-Quezada

pH Ecor (mV/SCE ) Icor (µA/cm2) Vcor ( mm/an)Passive domain

(mV)

UN TC UN TC UN TC UN TC

7.4 -110 -370 0.6 0.98 0.0061 0.0099 1210 1340

2.2 -160 -630 1.3 7.33 0.0131 0.0744 1079 1288

Contact angles decrease from 63.5 to 41.5 associated with a more hydroxiyl rich surface and with a cell viability 88.5%

Increasing and /or decreasing corrosion resistance using HA deposition

Electrode solution Vcor(mm/year)Stern-Geary Vcorr(mm/year)

Ion releaseTi Lactic acid 10% 0.048 0.056

Ringer 2 0.00016 0.00028Ringer2 + lactic acid 0.079 0.082

Ti+HA Ringer 2 0.121 -Ringer2 + lactic acid 0.0283 -

TiAlZr coated with HA

•

SEM image (a) and EDS spectrum (b) of HA/TiAlZrAfter puls electrodeposition without AgNO3 the TiAlZr surface area was completely covered by a thick Ca–P layer exhibiting the tell tale rose-like morphology observed as shown in Fig.a.

Chemical elemental analysis of the HA coating detected O,Al, P, Ca, and Ti as the main constituent elements, indicating that the coating is complex deposit on the alloy surface. The Ca/P ratio obtained on the surface was 1.64, almost the same value as the one of the hydroxyapatite (1.67), confirming the formation of the hydroxyapatite.

ab

TiAlVZr coated with Ag-HA

.The morphology of the HA/Ag coatings was revealed by SEM micrographs EDS line profile analysis was conducted in order to investigate Ag distribution in the film. The results proved that Ag was uniformly distributed in the film. The chemical elements of the Ag–HA film deposited on titanium were Ca, P, Ti, O and Ag. The Ca/P ratio obtained on the surface was 1.62.

TEM micrograph of silver nanoparticles loaded on hydroxyapatite

TEM images show a size distribution of globular-shape silver particles that range between less than 10 nm and 60 nm. It can be seen that the Ag particles preferentially adhere to the surfaces of HA. Dark little spots corresponds to Ag nanoparticles and larger spots corresponds to HA.

The density of attached nanocrystals is high. Observed Ag particles appear to have a narrow size distribution, and no free particles are observed in the background of the TEM images, which confirms all formed Ag nanoparticles are durably attached to the HA.

.

Electrochemical parameters for TiO2/TiAlZr, HA/TiAlZr, Ag-HA/TiAlZr

Electrochemical behaviorThe alloy covered with Ag-HA/TiAlZr exhibited lower corrosion rates than the alloy covered with HA, denoting the better efficiency of the Ag-HA coating in Hank solution. According to table with electrochemical parameters (where Ecor is corrosion potential, Icor, is corrosion current density and Rp polarization resistance) this behavior is sustained firstly by Ecor which is shifted to more electropositive values in this case. The corrosion rate for Ag-HA/TiAlZr coated specimen is 28.9 times lower then TiO2/TiAlZr and such values support the better anticorrosive properties of coating.

SamplesEcor (mV) Icor (μA/cm2) Rp(KΩcm2) Vcor (μm/y)

TiO2/TiAlZr -307 25.75 127.08 497.53

HA/TiAlZr -246 15.68 76.56 308.02

Ag-HA/TiAlZr -67 0.88 26.44 17.17

Antibacterial activity

For antibacterial activity evaluation three sterile samples were used as following:Escherichia coli (K 12-MG1655) was aerobically cultured in a tube containing Luria Bertani (LB) medium at 37 oC. The initial concentration of bacteria was adjusted to 10-9 colony forming units (CFU)/mL by dilution with sterile water.A quantitative analysis using 10-9 CFU of a gram negative bacterium which were cultured on LB agar plates containing: TiO2/TiAlZr, HA/TiAlZr; Ag-HA/TiAlZrAfter 24 hours, the presence of TiAlZr/TiO2 inhibited the bacterial growth by 22% (560·103 CFU/mL). In the presence of TiAlZr/HA was counted 250·103 CFU/mL with a inhibition a bacterial growth by 65%, while presence of Ag-Ha/TiAlZr caused 92% inhibition of bacterial growth (56·103 CFU/mL). For Control plate was counted 720·103 CFU/mL. According to above data bacterial ratio is increasing as following TiO2/TiAlZr, HA /TiAlZr Ag-HA/TiAlZr

Formation of bioactive surfaces by physiological methodThis method involves the heterogeneous nucleation and growth of bone-like crystals on the surface of implant at physiological temperatures and under specific pH.

To enhance the heterogeneous nucleation of the Ca-P on the titanium implants are used high concentration of Ca and PO43- is used in an increasing pH solution to form a thin layer on titanium surface

In electrolytes containing Ca2+ ions (without PO43-), small quantities of calcium are absorbed on the surface of the titanium oxide layer after 24 hours of immersion. This shows that the presence of phosphate ions is notmandatory for the Ca2+ ions to be adsorbed on the oxide surface.

The absorption of Ca2+ on the surface is most likely due probably to the electrostatic interaction of Ca2+ ions with the negatively charged surface.In physiological fluids containing phosphate and calcium ions, phosphate is absorbed on the surface of theTiO2, replacing hydroxyl groups OH− attached to the titanium ion as H2PO4

- − and/or HPO42-

H2PO4- ↔ HPO42- + H+

HPO42-↔ PO4

3- + H+

thus forming a strong and complex link.

Ca 2+ +HPO42- +2H2O ↔ CaHPO4 +2H2 O

8Ca 2+ + 6HPO42- +5H2O ↔ Ca8 H2 (HPO4)6 5H2O +4H+

3Ca 2+ +2PO43 +nH2 O ↔ Ca3 (PO4)2.nH2 O

10Ca2+ +6PO43- +2OH- ↔ Ca10 (PO4)6 (OH)2

The implant (TiAlMoFe) is first immersed in five-times concentrated SBF for 24 h at 37°C. A thin (<3 μm thick), dense and amorphous uniformly deposited layer of calcium phosphate upon the implant surface serves as a seeding substratum for the subsequent growth of a substantial (30 to 50 μm-thick) crystalline latticework.

After this first step the implants are immersed in Kokubo-SBF for various period of time. The initiations of growth of HA on the TiAlVMoFe support which in the initial stage presents a structure of dandelion bud that increases like an inflorescence. The Ca/P ratio obtained by EDS was 1.33 specific for octacalciumphosphate which is a precursor of the hydroxyapatite due to its high similarity with crystals present in bone and teeth (Leng et al. have identified such phase as OCP. According to Liu et al., OCP crystal assemblies seem to enhance osteocalcin expression, which is an important marker of osteoblast phenotype.

Biomimetic deposition of apatite on alkaline and heat treated titanium alloy surface.• Alkaline treatment is conducted to create bioactive layers on titanium surfaces by immersing them in thermal NaOH solution with subsequent heat treatment. The surface layer consists of an irregular and porous sodium titanate and TiO2 (rutile) Alkaline treatment generally includes immersion in 5-10 M NaOH at 60 °C for 24 hours followed by heat treatment at 600 or 800 °C for 1 hour. •Alkali and subsequent heat treatment probably causes low adhesive strength at the interface of the treated layer (about 1 μm in thickness) and the substrate change by distinct structures and compositions.

. The SEM photograph revealed that the control specimen had a smooth surface texture with abrasive marks

SEM photographs of the surfaces of the NaOH-treated Ti6Al4V alloy subjected to 5 and 10 M NaOH treatment at 60°C for 1 day

In contrast with control surface, the alkaline-treated specimen had porous surfaces. At the same alkaline-treating temperature, a much porous structure was observed with increasing the concentration of NaOH

The alkaline- and heat-treated titanium alloys leads to the formation of more porous surfaces. At the same alkaline-treating temperature, a more porous structure was observed with increasing treatment timeThe structural change on the titanium surface during alkali and heat treatments and mechanism of apatite formation on the treated surface in simulated body fluids are described as follows. During the alkali treatment, the TiO2 layer partially dissolves in the alkaline solution because of the attack by hydroxyl groups.

TiO2 +NaOH ↔HTiO3- + Na+

This reaction is assumed to proceed simultaneously with hydration of titanium.

43

2223

3

Ti(OH)HOTi(OH)

H21OHTiOeTi(OH)

4eTi(OH)3HOTi

↔+

+⋅→+

+→+

−+

+

+−

A further hydroxyl attack on the hydrated TiO2 produces negatively charged hydrates on the surfaces of the substrates as follows:

These negatively charged species combine with the alkali ions in the aqueous solution to produce an alkalinic titanate hydrogel layer. During heat treatment, the hydrogel layer is dehydrated and densifies to form a stable amorphous or crystalline alkali titanate layer.Coating of the Ti6Al4V substrates was performed by immersion of the disk-shaped samples in SBF at 370C for different periods of time. After immersion, apatite was formed on the surface of the specimens treated with 5 M NaOH at 80°C for 3 days when soaked in SBF for 1 day.

OnHHTiOHOOHnTiO 2322 ⋅↔+⋅ −−

• The biological activity of the interface is ensured by the ionic exchange, between the ceramic system (phosphate layers or bioglasses) and the physiological fluids. In the physiological fluids, ionic exchange surmises the elimination of sodium ions Na+ from glass or the layer of sodium titanates and their replacement with H3O+ ions produced in the surrounding fluid resulting in Ti–OH layer formation • Simultaneously, with increasing pH the apatite nucleationgets accelerated on increasing the supersaturation of thesolution with respect to apatite.• Calcium ions are incorporated in the hydrated Ti–OH layer. The positively charged Ca 2+

acts as nucleation sites for carbonate–hydroxyapatite attaching themselvesto the negatively charged (PO3

4- ) and (CO3

-2) to form Ca–P enriched surface layer which crystallizes to bone-like apatite (carbonate–hydroxyapatite)

The FTIR result of HA powder, alkali-heat treated TiAlV and alkali-heat treated TiAlV immersed in SBF

Formation of bioactive surfaces by electrochemical (HA) deposition• The electrochemical deposition of HA coating is normally performed in an aqueous solution containing calcium and phosphorus species (Ca–P). The conventionally used Ca–P concentrations are in the range of 10−1 to 10−3 M which are higher than or similar to that of human blood plasma. • The electrolyte used for obtained a HA coating on TiAlVZr has a molar ratio of Ca to P of 1.67. Cathodic polarization was conducted from open circuit potential to -3V (vs. SCE) at a rate of -0.6V/h. Coating were deposited at current densities of 1-20 mA/cm2 for 50 min at room temperature and pH 4. Through the electrochemical deposition process of calcium carbonates, uniform and adherent bioactive layers are obtained. The shape of the cathodic deposition curves is presented

Ca 2+ ions migrate to the cathode where the deposition occurs and can react with ions and OH- formed on the surface, thus synthesizing the hydroxiapatite deposition. The reactions which underlie the synthesis are:

2 34 10 4 6 2

2 34 2 3 4 2 2

2 24 2 8 2 4 6 2

2 24 2 4 2

10Ca 6PO 2OH Ca (PO ) (OH)

3Ca 2PO H O Ca (PO ) H O

8Ca 6HPO 5H O Ca H (PO ) 5H O 4H

Ca 6HPO 2H O CaHPO 2H O

n n

+ − −

+ −

+ − +

+ −

+ + ↔

+ + ↔ ⋅

+ + ↔ ⋅ +

+ + ↔ +

A SEM image at covered alloy: HA/Ti6Al4V1Zr is presented in Figure associeted with EDS spectrum. SEM micrographs revealed HA coatings to be entirely composed of straight plate-like units with sharp edges. Chemical elemental analysis by EDX of the HA coating detected oxygen, aluminium, phosphor, calcium, and titanium as the main constituent elements, indicating that the coating is complex deposit on the alloy surface. The Ca/P ratio obtained on the surface was 1.62,

Electrochemically coated HA surface produced highest level of proliferation of cell compared with uncoated surface after incubation 3 days on human osteoblast-like cells

HA coating on titanium with nanotubular TiO2 intermediate layer After the titanium metal was anodically oxidized in 0.5% HF + 5 g/l Na2HPO4, for 2h at 20V and subjected to heat treatment at 60º C for 1 h, could induce apatite formation in SBF because the amount of anatase and/or rutile was increased. In simulated body fluids, the titanium anodically oxidized inducing apatite formation on its surface, as shown The induction period of apatite formation decreased with increasing amount of either the anatase or rutile phase. It is well established that anatase phase is more efficient in nucleation and growth of apatite than the rutile phase of TiO2 because of the better crystal lattice match with HA phaseThe overall reaction for anodic oxidation of titanium can be representedas:At the Ti/Ti oxide interface:

At the Ti oxide/electrolyte interface: Ti 2+ +2O 2-↔TiO2 + 2e-

−+ +↔ e2TiTi 2

−+

+−

++↔

+↔

e4H4OOH2

H4O2OH2

22

22

Bioceramic coating of HA on titanium substrate with Nd-YAG laser

Schematic representation of monolayer HAp ceramic coating on Ti substrates. (a) Laser coating of HAp powders. (b) Composite microstructure Monolayer porous coating after laser coating.

Cross-sectional SEM micrograph of HAp coated Ti6Al 4V substrate using an Nd:YAG laser with power of 100W and scanning velocity of 1 mm/s.

SEM micrograph and chemical analysis by EDS on the cross-section of coating (a) HAp/Ti composites at interface, (b) HAp/Ti composites with depth increasing, (c) Titanium substrate.

The elastic constant results of the HAp coating on Titanium substrate from nanoindentation test

Schematic for the formation and morphology evolution mechanism of Ca5(PO4)3OH samples with various morphologies based upon different pH values

Acta Biomaterialiavolume 7, Issue 7, July 2011,

Nanoscale HA particles for bone tissue engineering

Schematic illustration of the formation of HAp nanorods.

Double-hydrophilic block copolymer (DHBC) poly-(vinylpyrrolidone)-b- poly(vinylpyrrolidone-alt-maleic anhydride)-b-poly-(vinylpyrrolidone) (PVP-b-P(NVP-alt-MAn)-b-PVP) was synthesised as the biomimetic template for the hydroxyapatite (HAP) nanocrystal synthesis. Needle-like HAP nanocrystals can be formed in the presence of PVP108-P(NVP-MAn)28-PVP108. TEM images of HAp nanocrystals

templated by PVP108–P(NVP-MAn)28–PVP108 after the 1 day (a), 8 day (b) and 13 day (c) preparations.

SEM micrographs of scaffolds fabricated with different polymer matrices.

Bone tissue cross-section of N-HA scaffold under phase contrast (a) 3 and (c) 12 weeks post-surgery with corresponding cross-polarized light micrographs shown in (b,d) representing birefringence of collagen strands. 200×original magnification. S, scaffold; M, mineralized bone; C, collagen birefringence; V, vessel.

Scheme of biomineralization mechanism of the preparation of HAp–PEG–MWCNTs

SEM images of the surfaces of the nAg–HA/TiO2 film and (B) SEM image ofTiAlZr covered with Tobrex–nAg–HA/TiAlZr,

• Morphology of the deposited films exhibited a plate-like structure with a small rod shape primary Ag nparticle on the films. The nAg–HA/TiO2 coated plate exhibited a number of global microstructures with diameters of 70–150 nm.

• These results are in good agreement with the DLS observation. Micro pores are detected on the thin film of nAg–HA/TiAlZrwith a diameter of 1–5 μm.

• The chemical elements of the nAg–HA/TiAlZr filmdeposited on TiAlZr alloys were Ca, P, Ti, Al,O Ag .

• There was 2.29 mass % of Ag in the material and Ca/P ratio was 1.61. The slightly lower ratio of Ca/P is probably caused by the substitution of Ag + ions for Ca2+ ions.

The tobramycin presents several hydrophilic groups, which could interact with ions present in the coating solution (i.e. Ca +2, HPO4 2− ).

Culture of E. coli inhibited by (A) TiAlZr uncoated (B) nAg–HA/TiAlZr (C) TiO2/TiAlZr (D) Tobrex–nAg–HA/TiAlZr

The antibacterial effect of biomimetic coating with silver nanoparticles is high and close to value of biomimetic coating with silver and antibiotic (having almost the same increase of antibacterial effect of 2.07 and 2.10 times).

Apatite-formation ability – Predictor of ‘‘bioactivity”?

• The ability to trigger the formation of apatite from a supersaturated solution has been widely used to imply the bioactivity of an implant in vivo. • The method itself may provide at best incomplete information, primarily because it is determined only by solution supersaturation, irrespective of biological processes. • Bone regeneration is triggered mainly by the vitality of osteoblasts, and regulated by the expression of growth factors such as oestrogen, parathyroid hormone and bone morphogenetic proteins,while ions or other species released from an implant may affect the expression of such growth factors,and so bone resorption or formation. • The misinterpretation of the outcome of such tests must result inmisunderstanding of the true effects and behaviour of materials intended for use in embedded biological contexts. Thus, the underlying and motivating hypothesis needs to be carefully reconsidered, along with the results of all work founded on the concept. It would seem that it is only viable to test using osteoblasts,whether in vivo or in vitro.

Acta Biomaterialia 6 (2010) 4181–4188

• Since the 1970s, a bioactive glass, so-called 45S5 bioglass,has shown extraordinary performance in bonding spontaneously with living tissues, apparently through the formation of biological apatite (BAp) as an interfacial layer. Likewise, glass ceramics and wollastonite have shown similar behaviour in vivo. Furthermore,nearly complete conversion to apatite-like mineral was observed with the degradation of b-tricalcium phosphate (TCP) and CaSiO3 Given such observations, on rather chemically diverse substrates, ‘‘apatite-forming ability” (AFA) – that is, theoccurrence sooner or later of an apatitic layer on the implant –has been widely assumed to be a predictor of bioactivity in vivo.. •A recent online search found that more than 1100 articles have been published on this topic since 1990. However, this phenomenon may provide incomplete or misleading information. Indeed, as pointed by Bohner and Lemaitre the assumption of validity lacks support from ‘‘enough scientific data”. Even then,the proposal was only for a modified solution and test protocol.

•The key concern addressed here is that bioactivity cannot be predicted from observed AFA simply on immersion in a so-called ‘‘simulated body fluid” (SBF).

SBF was designed by Kokubo and only contained inorganic ions with concentrations similar to those of human extracellularComponent concentrations in blood plasma and various ‘‘physiological”solutions (mM).

Na+ K+ Mg 2+ Ca 2+ Cl- HCO3- HPO4 2- SO4-

Blood plasma 142. 3. 6–5.5 1. 2.1–2.6 95.–107. 27. 0.65–1.45 1.0Human blood plasma 142. 5.0 2.5 1 103 . 27 1.0 0.5Original SBF 142. 5.0 2.5 1 148 4.2 1.0 0.Corrected SBF (c-SBF) 142. 5.0 2.5 1 147 4.2 1.0 0.5Revised SBF) 142. 5.0 1.5 2.5 103. 27. 1.0 0.5Ionized SBF 142. 5.0 1.0 1.6 103. 27.0 1.0 0.5Modified SBF 142. 5.0 1.5 2.5 103.0 10.0 1.0 0.5Newly improved SBF 142.0 5.0 1.5 2.5 103.0 4.2 1.0 0.5

• later the role of bicarbonate HCO3- was emphasized ;the original solution wasapparently meant to be comparable with human extracellular fluid on the basis of the total analytical concentrations of key components.

• It was then remarked that, at physiological pH (7.40), the ion activity product (IP) for such solutions with respect to [Ca2+], [PO4

3-] and [OH-] is much greater than corresponds to saturation for hydroxyapatite (HAp) at that same pH. • the calculated degree of supersaturation, S= (IP/Ksp ) 1/n where Ksp is the solubility product and n is the number of ions in a formula unit. They gave S= 12

•Hence, the creation of a solution to mimic tissue fluid is assumed to require such analytical concentrations.This is overly simplistic in that it fails to recognize that tissue fluid contains many substances, including proteins, that bind to both Ca2+ and PO4

3- ions such that the actual ion product is substantially affected.• However, to focus on this alone is to miss the point: normal tissue fluid cannot in fact be supersaturated with respect to apatite or any calcium phosphate.

• Saliva has long been said to be supersaturated with respect to tooth mineral –HAp But, the role of binding to protein has been completely ignored. If left standing, saliva produces precipitates of calcium phosphates. This has been used as supposed evidence for HA supersaturation, and attributed to the rise in pH due to loss of CO2, with bicarbonate being conventionally described as the principal buffer system of saliva.

•However, the complexation of calcium by carbonate with phosphate has not been recognized, even though the chemistry has long been known .Such complexes can account for a substantial proportion of the analytical concentration of both Ca 2+ and PO4

3-. Indeed, there are many other ligands which must be taken into account in the speciation of such complicated solutions, and in the absence of the detailed calculation a statement regarding degree of saturation is not meaningful.

•Indeed, it appears that many solution species and equilibria may not be documented, making the calculations even more difficult.

•Recent work has also shown that the solubility of HAp is not as expected, but substantially lower than conventionally believed, even in a very simple solution

Trace elements• Although the biological responses to trace elements in the human body are not yet fully understood, such ions have already shown special roles in the maintenance of bone health or the stimulation of bone regeneration, in particular Zn, Sr, Si, and recently B. Cells are necessarily sensitive to the local environment; such ions affect not only the activity of osteoblasts, butalso apatite nucleation. For example, Sr has been found to be effective in the treatment of osteoporosis by enhancing pre-osteoblast differentiation, inhibiting osteoclast differentiation and reducing osteoclast function. It has therefore not only been suggested as a daily oral drug, but has also been incorporated into Hap, bioglass, , b-TCP and CaSiO3.

• Sr has been reported to be selectively distributed, with 3–4 times more in new than in old cortical bone, and approximately 2˝ times more in new than in old cancellous bone. The formation of new bone may involve the preferential precipitation of partially Sr-substituted apatite (Sr-HAp) during Sr treatment, although it is not a normal constituent of natural human bone (typically, about 320 mg in total in the human body, mostly in bone and connective tissue)

Argument crystallographic. •‘‘Apatites”,, has a remarkable structural versatility, with variations in both lattice and morphology (crystal habit), and suffering ready solid-solution substitutions by many foreign ions.

•However, it has yet to be resolved just what kind of apatite structure is favoured for bone mineral. Apatite may be represented as Ca5(PO4)3X, where X = OH, F or Cl (at least, these are the ions of major biological interest). The major, and therefore the most important, biomineral is an impure form of HAp having a hexagonal structure, sometimes referred to as ‘‘biological apatite” (BAp)• In contrast to synthetic stoichiometric HAp, Bap is a calcium-deficient apatite with various random substitutions at the HAp lattice sites by common ions such as Na+, K+ andMg2+, which partially substitute for Ca 2+, while CO3

2-, Cl- and F-may occupy OH- sites. A range of trace elements, such as Sr 2+, Pb4+, Ba 2+, Fe 3+, Zn 2+ and Cu 2+, may substitute at Ca 2+ sites,while various other substitutions are feasible at PO4

3-, sites, especially HPO42- when there is an overall cation deficiency. Such substitutions lead to changes in the values of lattice parameters and distortion of the crystal structure Compared with wellcrystallized synthetic HAp, BAp is poorly crystalline. While tooth enamel mineral shows better crystallinity, dentine and bone are the least-well crystallized, and may even (appear to) be relatively amorphous. ‘‘Apatite” must be treated as a structure type rather than a specific compound.

•Bone is formed by the mineralization of an organic matrix (largely collagen), through the nucleation and growth of a mineral closely similar to HAp, regulated by osteoblasts – multinucleated bone-formation cells, which are activated in an alkaline environment (Fig. ) – which have the capacity to proliferate, to synthesize organic bone matrix and to respond to growth factors. •It follows, that bioactivity expressed simply by apatite forming ability under artificial and non-biological circumstances is meaning less as a predictor, possibly despite such a process being observed in the presence of osteoblasts (Fig.).

•If an apatitic layer is first formed, it does not follow that the surfacewill necessarily be colonized by osteoblasts or, if it is, that they willnecessarily form bone. In any case, that layer is likely to be resorbedsooner or later, even if it is conducive to subsequent cell and bone growth, and therefore implant success. •it does appear that pre-existing apatitic coatings are conducive to the development of adherent bone, and the current intense researchactivity in this direction is justified.

•it has never actually been demonstrated that the first and crucial step in the biological process of ‘‘accepting” an implant in vivo is the spontaneous precipitation of any apatite

•let alone a BAp, on its surface as an essential precursor to cell contactand growth.

•A novel anodic TiO2 nanostructural layer was reported composed of nano-channels and showing a much higher mechanical stability. This mesoporous titania layer (MTL) reportedly has a higher surface area and better adheres to the Ti substrate under stress.

•The bioactivity of this nano-channel surface was investigated in the light of itsHAp formation ability. Apart from the kinetics, the structure andchemical composition of HAp obtained from two standard SBFs were also compared with the results from the (up to now) most efficient nanotubular system.To prepare an approximately 10 lm thick MTL structure, anodization of the Ti foil was carried out at 1 V in dehydrated glycerol/K2HPO4 solution at 180 C for 3 h in a two electrode system The samples then were cleaned by overnight washing in distilled water and dried in a N2 stream. For etching treatments the anodized samples were sonicated in 30 wt.% H2O2 in a sonication bath for 1 h at room temperature. Annealing of samples was carried out in a Rapid Annealer at 450 C in air at a heating and cooling rate of 30 C min1 convert the as-formed layer into anatase/rutile phase.

Acta Biomaterialia xxx (2011) xxx–xxx

In order to investigate the formation of apatite on the different TiO2 samples we used two different compositions of SBF: (i) the classic Kokubo solution and (ii) a modified solution proposed by Bohner and Lemaitre.

Briefly the Kokubo (KSBF) solution with ionic concentrations equal to human blood plasma was prepared by sequentially dissolving reagent grade chemicals, including NaCl, KCl, NaHCO3, MgSO412H2O, CaCl2, and KH2PO4, in double distilled water at 37 C and buffered with TRIS and HCl at pH 7.4. The chemical composition is as follows (mmol l1-): Na+,142.0; K+, 5.0; Mg 2+ , 1.0; Ca 2+, 2.5; Cl-, 131.0; HCO3

- , 5.0; HPO42- 1.0; SO42- 2, 1.0

The Bohner and Lemaitre solution (cSBF3) was prepared as in Bohner and Lemaitre paper.

To induce HAp growth 1 cm2 samples were soaked in the 50 ml SBF solutionunder static conditions in a biological thermostat at 37 C for different days. After the experiment the samples were rinsed with 20 ml of deionized water and dried at 60 C for 1 h.

SEM images of cross-sectional (inset: top) views of (a) the as-formed (non-etched) mesoporous titania (MTL) layer showing the presence of a thin top layer the entire oxide layer after formation consists of interlinked regular channels of 8–12 nm diameter. However, for the ‘‘as-formed’’ layer a thin, dense surface layer is present on top of the open channels(b) The MTL layer after etching in H2O2solution. (c) XRD patterns of the MTL layers after

different annealing and etching treatments. Am-NE, amorphous non-etched; Am-E,amorphous etched; An-NE, annealed non-etched; An-E, annealed etchedthe top layer had been removed and the channels were approximately 12 nm in diameter. XRD investigations of the MTLbefore and after etching show the layers to be mostly amorphous, but they contain some anatase and rutile phase

•SEM images of HAp formation on different MTL surfaces as well as TiO2 nanotube surfaces for different immersion times in (a) cSBF3 and (b) KSBF solutions. Am-NE, amorphous non-etched; Am-E, amorphous etched; An-NE, annealed non-etched; An-E, annealed etched; NT nanotubes.

•The MTL layers (after etching and etched &annealed) show the highest coverage for the cSBF3 solution (a), with a complete 1µm thick layer after 6 days. The amorphous non-etched and only etched MTL layers show significantly less HAp formation after 2 and 6 days. •The best surface, etched anatase MTL, shows HAp formation over the entire surface after 2 days. In comparison, on anatase nanotubes, HAp formation could be observed after 2 days, but thickness of HA is only 100 nm after 6 days.If KSBF solution is used (2b) HAp formation can only be observed in the case of etched MTL samples (amorphous as well as anatase) after 4 days and the highest growth rate is clearly obtained on the etched and annealed surface.

• In comparison with cSBF3, KSBF leads to much slower HAp formation and after 14 days a thick layer covering the surface for the MTL. The etched anatase MTL layer shows better coverage by HAp compared with an etched amorphous MTL layer.

• Morphologically, HAp formation in the two different solutions is different. In KSBF solution globular HAp was observed on MTL, forming large plates over the entire surface after 14 days incubation. This visibly resembles previously reported structures of crystalline HAp on nanotube surfaces

• In contrast, in cSBF3 a rather spongy morphology of the HAp precipitate can be observed, which is intuitively in line with very fast HAp formation.General trend of Hap formation remains the same, i.e.: •The etched,annealed MTL surface always shows the most efficient HAp formation• Mass gain of the samples after HAp loading for different durations in cSBF3 and KSBF solutions.( Am-NE, amorphous non-etched; Am-Et, amorphous etched; An- NE, annealed non-etched; An-Et, annealed etched; NT, nanotubes); confirm the SEM observations, i.e. that the etched anatase MTL layers are the most efficient surface for HAp formation in comparisonnot only with other forms of MTL but also compared with nanotubularsurfaces.

Comparison of in-growth HAp formation in etched anatase MTL layers: cross-sectional SEM images of MTL layers (a) before and (b and c) after soaking in (b) KSBF and (c) cSBF3 after 6 and 14 days, respectively. For KSBF the pores are filled with HAp (complete in-growth); for cSBF3 only a top layer is formed.

•For HAp layers formed from both the cSBF3 and KSBF after 6 and 14 days, XPS analysis reveals a similarcomposition.

•Both apatite layers show traces of Mg and Na in addition to the main components Ca and P. Mg 2+ and Na+

ions might substitute Ca 2+ in apatites,in line with the composition of biological apatites from human bones, (non-stoichiometric).

•Traces of ions Mg 2+ and Na+, as well as the anions CO3

2-, HPO42-, F-,and Cl-,

are also present in natural apatite. A qualitative analysis shows that the composition of the apatite layers, the Ca/P ratio, are close to physiological Hap, which has a Ca/P of 1.35–1.46 ;apatite formed from KSBF has lower Ca /P

•The nature of the apatite forming substrate, amorphous or annealed, or the morphology, channel or tubular structure, do not seem to influence the final composition of the apatite.

• XRD patterns of the HAp layers on the etched anatase MTL surface after 6 and 14 days soaking in cSBF3 and KSBF.• The XRD pattern of the HAp layer grown in cSBF3 solution shows peaks corresponding to anatase TiO2 as wellas a small peak at 32 for HAp, whereas the HAp layer grown in KSBF solution only shows peaks of HAp. This difference is due to the morphological difference in HAp formation in the two different solutions as well as the in-growth phenomena of Hap• The magnified XRD pattern at 32 shows that both HAp layers are similar in terms of crystallinity.• However, in case of cSBF3 solution this weak peak overlaps with the sharp anatase peak from the substrate.