Materials Science and Engineering C 32 (2012) 321–329
Contents lists available at SciVerse ScienceDirect
Materials Science and Engineering C
j ourna l homepage: www.e lsev ie r .com/ locate /msec
Zr–Ti–Nb porous alloys for biomedical application
A.E. Aguilar Maya a, D.R. Grana b, A. Hazarabedian a, G.A. Kokubu b, M.I. Luppo b,⁎, G. Vigna b,⁎a Materials Department, Argentina Atomic Energy Commission, Gral Paz 1499 (B1650KNA), San Martín, Buenos Aires, Argentinab Anatomical Pathology, School of Dentistry, Faculty of Medicine, Argentina Association of Odontology, University of El Salvador, Tucumán 1845 (1050), Buenos Aires, Argentina
Recent studies linked to the production of implants focus on the development of porous materials, as theyprovide good biological fixation to the surrounding tissue through bone tissue ingrowth into the porous net-work.Research on the biological behavior of metals has shown that the composition of implant biomaterials mustbe carefully selected to avoid adverse reactions. Ti, Zr and Nb are non-toxic metals with a good compatibility.In the present study, Zr–Ti–Nb foams of two compositions (Zr–34.4%Ti–1.6%Nb and Zr–34.5%Ti–5.5%Nb)were fabricated starting from hydride–dehydride powdered metal using space-holding fillers. Both foamsdisplayed an interconnected porous structure with a porosity of 70%. The average pore size was around260 μm. The Young's modulus and the compressive plateau stress were observed to vary with the Nb contentin the range of 0.3–1.4 GPa and 11–32 MPa, respectively. All alloys tested – in porous and solid forms –
showed excellent biocompatibility in subcutaneous as well as in bone tissues. The alloy with more Nb contentshowed pronounced osteoinductive properties.
Materials for bone replacement might mimic the architecture ofbone [1]. Most implants are metals: stainless steels, Co–Cr system al-loys and titanium alloys [2]. The Young's modulus of those biomate-rials is much greater than that of cortical bone, then boneresorption occurs [2]. That is because when the bone is stressed, thebone-producing cells called osteoblasts are stimulated into generat-ing more bone. So if the bone is replaced by a metallic counterpartthat is stiffer than the original bone, the replacement will tend tobear a greater proportion of the load, shielding the surrounding skel-eton from its normal stress levels. Thus, bone replacements cementedto neighboring bone may become loose over time, as the surroundingbone is resorbed. For this reason, the life time of metal bone replace-ments is not longer than ten years.
The ideal bone substitute is not a material that interacts as little aspossible with the surrounding tissues, but one that will form a securebond with the tissues by allowing new cells to grow. One way ofachieving this is to use a porous material, so that new tissue, and ul-timately new bone, can grow into the pores and help to prevent loos-ening and movement of the implant. Recently, a new series of highlyporous metals has been developed and released for use in orthopedicsurgery: porous tantalum and three porous titanium foams [3].
rgentina Atomic Energy Com-os Aires, Argentina. Tel.: +54
Numerous researchers conclude that a porous matrix may be usedas implant if it meets with the following requirements [4,5]
• Interconnected porous structure to provide necessary space for cellingrowth and vascularization.
• High porosity and pore size in the range of 200 to 500 μm; other-wise, the new tissue cannot develop an effective blood supply.
• The elastic modulus of bone ranges from 0.1 to 20 GPa.• The compressive strength ranges from 2 to 200 MPa• Alloys composed of non-toxic and non-allergenic elements.
As reviewed by Banhart [6] there are many ways to manufactureporous metallic structures and they can be classified according tothe state the metal is processed in. One possibility is to start fromsolid metal in powdered form. There are various technologies to pro-duce powder of IVb group elements where the hydride–dehydride(HDH) process is among the most common procedures [7]. In theHDH process, high quantities of hydrogen are introduced at elevatedtemperatures. During cooling, brittle hydrides are formed. Then, thematerial can be easily milled at room temperature. Heating of this hy-dride powder in high vacuum allows hydrogen desorption thus pro-ducing the metallic powder. Metallic foams can be obtained bymixing filler materials and metal powders. A suitable solvent oreven an organic binder may be used to mix the space holders andthe metal powders. Ceramic particles or hollow spheres, polymersgrains or hollow polymer spheres, salts or even metals can be usedas space holders. The mixture is followed by cold pressing, removingthe space holders by thermal treatment, and finally by sintering [6].
322 A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
The selection of Ti and Ti alloys for implantation is determined bya combination of favorable characteristics including immunity to cor-rosion, biocompatibility, osseointegration, excellent strength-to-weight ratio, good fracture toughness, high fatigue strength, andlow modulus and density [2,8–11]. These characteristics increasethe use of titanium alloys as biomaterials when compared to moreconventional stainless steels and cobalt-based alloys.
The properties of Ti alloys are sensitive to their phases/crystalstructure, and certain phases may be stabilized by the addition ofalloying elements [12]. A good candidate for alloying is Zr, which is aneutral element when dissolved in Ti and it can enhance strengthand improve elasticity of alloys [13]. On the other hand, Zr is amaterialof interest for surgical implants because it shows acceptable mechan-ical strength, satisfactory biocompatibility, good osseointegration,good corrosion resistance [14]. More studies comparing Zr and Ti im-plants showed that the degree of bone–implant contact is actuallyhigher in the case of Zr [15]. Both Zr and Ti belong to the same groupin the periodic table of elements and the Ti–Zr system shows as a com-plete solid solution [8].
Alloys containing elements such as Nb, Zr, Ta, Pt and Ti are beingextensively evaluated since these are the only five elements thathave been identified as producing no adverse tissue reaction[2,16,17]. In Ti alloys, Nb is a β-stabilizer [18]. The addition of Nbcould improve the strength keeping the Young's modulus low [19].In particular, Wang et al. [20] have fabricated a porous Ti–10Zr–10Nb alloy with a porosity of 69% which exhibits a Young's modulusof 3.9 GPa and a compressive plateau stress of 67 MPa, resemblingthemechanical properties of cortical bone. However, as bonemechan-ical properties are highly variable according to species, age, anatomi-cal site, liquid content, etc., once the porosity of the implant isselected, it is useful study the possibility to adjust the values of theYoung modulus and compressive plateau stress by changing theamount of a minority element, as Nb, in the Zr–Ti alloys given thatthey are promising biomaterials because of their biocompatibility.
It is of utmost importance that these materials have a biocompat-ible chemical composition to avoid adverse tissue reaction, which is afeature of clinical significance for materials implanted in long-termclinical situations in both human and veterinary medicine as therehave been some links between prolonged exposure to non-biocompatible materials and neoplastic tissue responses [21,22].
In the present study, Zr–Ti–Nb foams of two compositions werefabricated starting from HDH powdered metal using space-holdingfillers. The characteristics of the porous alloys, and their mechanicalproperties were evaluated. A comparative study of biocompatibilityof both Zr–Ti–Nb in solid and porous form was performed in Wistarrats' tibia bones and subcutaneous cellular tissue.
2. Experimental
2.1. Materials composition and foaming process
Two Zr–Ti alloys (see Table 1) with different Nb contents, wereprepared by remelting the starting alloys shown in Table 1 in an arcfurnace with high-purity Ar atmosphere, in order to obtain buttonsof approximately 20 g.
The hydriding process was perfomed in a Sieverts' apparatus [23],which consists of a quartz tube (25 mm diameter and 500 mm length)where the samples to be charged are introduced, one end of the tubebeing closed and the other end connected to a vacuum and gas input sys-tem. The heat treatments were performed by sliding an instrumentedfurnace on the hydrogenation chamber. Temperature was measured bymeans of a thermocouple installed in the chamber. Hydriding was doneat 600 °C and pressures close to 0.1 MPa (Hydrogen purity: 99.999%).
The hydrided samples were milled, under vacuum, with a RetschPlanetary ball mill PM400 model using stainless steel recipients and20 mm diameter balls, at 200 rpm during 10 min. The obtained
powder was separated via mechanical sieving in a vibrating sievewith aperture size of 37, 74 and 125 μm.
The powder, whose size was between 37 and 74 μm, was put in astainless steel semi-cylindrical capsule – covered with a mesh of 30 μm– during dehydriding to avoid being absorbed by the primary rotarypump in the former pumping stages. The dehydriding was carried outin the Sieverts' apparatus at 600 °C in dynamic vacuum conditions.After dehydriding, high purity argon was injected into the quartz tubeup to room atmosphere in order to prevent powder ignition. Finally, theargon was gradually replaced with air. During dehydriding, the powderwas partially sintered, and then it had to be milled and sieved again.
The metallic powders, with particle sizes between 37 and 74 μm,were mixed with ammonium hydrogen carbonate particles in a Y-type mixer. The ammonium hydrogen carbonate particles were actingas a space-holding material in this process. The particle size of theammonium hydrogen carbonate particles, between 300 and 700 μm,were obtained by sieving.
Cylindrical cross-section green compacts (diameter: 13 mm andlength: 26 mm) of the mixture were produced by uniaxial cold press-ing at 400 MPa.
The green compacts obtained were subjected to two heat treat-ments: the first at 200 °C for 5 h to burnout the space-holding particles,and the second at 1300 °C for 2 h to sinter the foams (diameter: 11 mmand length: 24 mm).
2.2. Foam characterization
The characterization of the porous alloys was carried out using opti-cal (OM) and scanning electron microscopes (SEM). In order to studytheir microstructures, the samples were etched in a 30 ml distilledwater, 60 ml H2O2 (100 vol.), 10 ml HF and 15 ml HNO3.
The pore size distribution was measured by quantitative imageanalysis. Previously, the foams were embedded in a black resin andpolished to prepare the plane for observation by optical microscopy.
X-ray diffraction (XRD) was performed to identify the phases pre-sent and lattice parameters of the samples. The diffractometer wasoperated at 45 kV, 40 mA, using CuKα radiation. In order to quantifythe phases present in the samples, X-ray diffraction patterns were an-alyzed with the PowderCell software using the structural data and theatomic positions collected from literature.
Densities were determined using a Micromeritics AccuPyc 1340helium pycnometer
Compression tests were carried out in a Shimadzu testing machineat room temperature with a strain rate of 10−3 s−1. The strain wasmeasured using an extensometer attached to the specimen. Threetests were performed on each type of alloy. The fracture surfaces wereevaluated using SEM. Due to the excessive shrinkage produced duringthe sintering, which distort the shape of foams, rectangular prism spec-imens for compression test were taken from the center of the sinteringfoamswith a length to diagonal of the base ratio of 1.5–2 (length~13 mmand diagonal of the base~8 mm).
Fig. 1. Optical micrograph of a) Nb1.6 and b) Nb5.5 alloys as remelted.
Table 2Crystalline structure and lattice parameters of the α″ phase in the TiZrNb alloys asremelted.
Phase Crystallinestructure(Space group)
Nb1.6 Nb5.5
Lattice parameters(nm)
Lattice parameters(nm)
α″ Orthorhombic(63)
a=0.3067b=0.5417c=0.4911
a=0.3058b=0.5414c=0.4903
323A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
2.3. Animal model
Thirty male Wistar rats (300–350 g) were included in this study.The animals were housed at controlled temperature and light–darkcycles of 12 h with ad libitum access to food and water. The ratswere handled and maintained in accordance with international rec-ommendations [24].
For biocompatibility evaluation the subcutaneous implantationand the tibiae implantation tests were performed. In this regard, therats were divided in 3 groups composed by 10 rats each one:
I. Materials Nb1.6 and Nb5.5 as remelted were implanted in leftand right tibias, respectively in 5 animals. In the rest the sameprocedure was performed but employing materials Nb1.6 andNb5.5 foams. Animalswere sacrificed at 7 days following surgery.Samples were prepared undecalcified (see below).
II. Same as I but animals were sacrificed at 60 days following sur-gery. Samples were prepared undecalcified.
III. Same as I, but materials were also implanted in subcutaneouscellular tissue. Animals were sacrificed at 7 days and sampleswere prepared decalcified.
All surgery was performed under sterile conditions. General anes-thesia was obtained by intraperitoneal administration of ketaminechlorhydrate (14 mg kg−1) and acepromazine (10 mg kg−1).
2.3.1. Subcutaneous implantationThe dorsal skin was shaved and disinfected with povidone–iodine
solution and two incisions (approx. 15-mm long) were made using ascalpel. By blunt dissection 2 pockets were prepared. Into each one, adisk either of Nb1.6 as remelted or foam and Nb5.5 as remelted orfoam was implanted. Each material was, at least 20 mm apart fromthe incision line into the subcutaneous tissue to avoid interferenceof the tissue response to each material. Finally, the wounds wereclosed with cyanometacrylate.
After sacrifice, at 7 days after surgery, subcutaneous tissue aroundimplantation areawas retrieved andfixed in 10% buffered formaldehyde.
2.3.2. Tibial implantationThe rats' legs were clipped of all hair, prepared with a povidone–
iodine solution and draped for surgery. A longitudinal skin incisionwas made in order to expose the tibial crest in both legs. After incisingand raising the periosteum, a cortical windowwas excised with a hol-low drill with sterile saline irrigation. In the left leg, a rectangularprism (width~height~1.5 mm and length~2 mm) of Nb1.6 asremelted or foam implant was placed inside the bone and the woundswere closed in layers with silk. In the left leg, a similar procedure wasperformed in order to place a rectangular prism (width~-height~1.5 mm and length~2 mm) of Nb5.5 as remelted or foam,depending on the study group.
This procedure resulted in 15 rats carrying Nb1.6 as remelted andfoam implants and another 15 animals with Nb5.5 as remelted andfoam implants. Neither antibiotics nor anti-inflammatory or analgesictherapy was employed. The rats were carefully supervised in thepost-surgical period for complications including pain, discomfort orinfection. Animal health was also monitored through changes inbody weight.
Healing periods of 7 days and 60 days were allowed depending ofthe assigned group. After 7 days and 60 days, the animals were killedwith an intravenous anesthetic overdose. Following euthanasia, thetibiae were removed and fixed in 10% buffered formol for 48 h.
2.3.3. Preparation of samplesUndecalcified sections (groups I and II) were embedded in methac-
rylate resin, ground and polished, and finally stained with aniline blue.
After fixation, subcutaneous (SC) samples and tibiae from group III(decalcified with EDTA) were included in paraffin and stained withhematoxylin-eosin.
The specimens were blindly evaluated by two independent ob-servers using a light microscope coupled with a digital cameraCanon Powershot A510 in order to study the bone-to-contact areas.
The retrieved micro-implants (group III) and the embedded sam-ples in methacrylate were observed under environmental scanningelectronic microscope.
3. Results and discussion
3.1. Material
Optical micrographs (Fig. 1) of the Zr–Ti–Nb alloys as remeltedshow the aspect of the α″ phase in these alloys as it was verified byX-ray diffraction (Table 2).
Fig. 2 shows the typical angular morphology of the Zr–Ti–Nb pow-ders obtained by HDH method.
Fig. 2. SEM micrograph of the Zr–Ti–Nb powders obtained by HDH method.
Fig. 3. SEM micrograph of the Zr–Ti–Nb foams: a) Nb1.6; b) Nb5.5. The β phase ispointed out with an arrow.
324 A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
The measured density and the calculated porosity of the Zr–Ti–Nballoys as remelted and the Zr–Ti–Nb foams are shown in Table 3. Theobtained densities for Zr–Ti–Nb foams were close to those measuredfor the Zr–Ti–Nb as remelted. Then, an interconnected porous struc-ture was obtained for both types of Zr–Ti–Nb foams and a high poros-ity – around 70% – was achieved.
The results obtained from the analysis of the X-ray diffraction pat-terns are shown in Table 4. A bimodal microstructure consisting of αand β phases is present in both foams, but at different volume frac-tions. Even though the ω phase – a brittle phase usually present inthese particular alloys – was taken into account, the only phasesfound were the presented in Tables 2 and 4.
The quantity of β phase (the lighter phase) increases with the Nbcontent, which can also be seen in the micrographs presented in Fig. 3.
Fig. 4 shows the surface of the Zr–Ti–Nb foams. Two types of poreswere observed: macropores and micropores. The macropores wereobtained as a result of the evaporation of ammonium hydrogen car-bonate particles and they are interconnected throughout the wholefoam. The micropores were obtained due to the incomplete sinteringof metallic powders [10] and/or from the volume shrinkages that oc-curred during the sintering process [4].
Fig. 5 shows the opticalmicrographs of the cross sections of Zr–Ti–Nbfoams embedded in a black resin and polished. From these micrographsthe macropores has been studied by means quantitative image analysis.
Table 3Densities and porosities of the TiZrNb alloys as remelted and the TiZrNb foams.
Nb1.6 Nb5.5
As remelted Foams As remelted Foams
Density(g/cm3)
Density(g/cm3)
Porosity(%)
Density(g/cm3)
Density(g/cm3)
Porosity(%)
5.6±0.1 5.3±0.1 71 5.6±0.1 5.4±0.1 70
Table 4Phases in the TiZrNb foams.
Nb1.6
Phase Crystalline structure(Space group)
Lattice parameters(nm)
α (TiZr) Hexagonal(194)
a=0.3099c=0.4917
β (TiZrNb) Cubic(229)
a=0.3461
Fig. 6 shows the histograms of macropore size distribution of theZr–Ti–Nb foams. The mean values, obtained from fitting with a Gauss-ian function, were 257 μm for the Nb1.6 foams and 269 μm for theNb5.5 foams. The size of macropores lay between 100 and 800 μmand they differ with regard to size distribution of the space-holderparticles due to the sintering shrinkage.
Fig. 7 shows typical compressive stress–strain curves obtained bytesting the Zr–Ti–Nb foams. They show linear elasticity at low stres-ses, where the initial slope is the Young's modulus, followed by along collapse plateau, truncated by a regime of densification inwhich the stress rises steeply [25]. This last stage was interrupted tosave the foams for fractographic observations.
The Young's modulus and compressive plateau stress for the Zr–Ti–Nb foams (Table 5) were determined from the stress–strain curves ofFig. 7.
Fig. 4. SEM surface topographical images of the Zr–Ti–Nb foams: a) Nb1.6; b) Nb5.5.
Fig. 5. Optical micrographs of the cross sections of Zr–Ti–Nb foams embedded in ablack resin and polished: a) Nb1.6; b) Nb5.5.
Fig. 6. Pore size distribution of the Zr–Ti–Nb foams fitted to a Gaussian function:a) Nb1.6; b) Nb5.5.
325A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
In order to evaluate the compositional effect on the mechanicalproperties it was necessary to compare alloys with the same porosity.Therefore, the Young's modulus and compressive plateau stress were“normalized” with respect to the porosity through the followingequations [25]
E P1ð ÞE P2ð Þ ¼
100−P1ð Þ100−P2ð Þ
� �2ð1Þ
σ P1ð Þσ P2ð Þ ¼
100−P1ð Þ100−P2ð Þ
� �2ð2Þ
where E(Pi) and σ(Pi) are the Young's modulus and compressive pla-teau stress for the Zr–Ti–Nb foams with a Pi porosity, respectively.
It is noted that once the porosity is fixed, the Young's modulus andcompressive plateau stress depend on the amount of Nb: the higherthe percentage of Nb the lower the mechanical properties. Then,once the porosity of the implant is selected, it would be possible toadjust the values of the Young modulus and compressive plateaustress by means of the Nb content in the alloy.
The pore size and mechanical properties of foams produced fromTi–Zr–Nb alloy are comparable to concurrent metallic foams for med-ical applications. To cite a few examples, Wen et al. [4] obtained Tifoams with a porosity of 78%, a compressive strength of 35 MPa anda Young's modulus of 5.3 GPa. These values were a close match tothose for cancellous bone which have a compressive stress between2 and 10 MPa and a Young's modulus between 1 and 10 GPa. Ti–Zr
Fig. 7. Nominal stress–nominal strain of the Zr–Ti–Nb foams under compressiveloading.
Fig. 8. SEM images of the fracture surfaces of the compression tested Zr–Ti–Nb foams:a) Nb1.6, b) Nb5.5.
326 A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
alloy foams with 70% porosity with a pore size ranged from 200 to500 μm showed a compressive plateau strength of 78.4 MPa and aYoung's modulus of 15.3 GPa [5]. Wang et al. [26] fabricated a Ti–10Nb–10Zr foam with macropores between 300 and 800 μm. TheYoung's modulus and the compressive strength with porosity of 69%were 3.9 GPa and 67 MPa, respectively. Kotan et al. [27] producedTi–6Al–4 V foams in the 60–70% porosity range with mean poroussizes of 400 μm and for the foam with a 70% porosity they obtainedan elastic modulus of 0.5 GPa and a compressive plateau strength of14 MPa.
A new series of highly porous metals has been developed and re-leased for use in orthopedic surgery all with a characteristic appear-ance similar to cancellous bone. Between them, Regenerex™ is ahighly porous titanium alloy with an average pore size of 300 μm,an overall porosity of 67% and a modulus of elasticity of 1.6 GPa.The clinical use of Regenerex™ was started at the beginning of 2007[28].
The fracture surfaces of the Zr–Ti–Nb foams exhibited both ductileand brittle fractures as it is shown in Fig. 8. Ductile dimples were pre-dominant in the Nb1.6 foams. Smaller areas of less dimpled regionswere observed in the Nb5.5 foams which tended to exhibit more fac-eted brittle fracture surfaces.
3.2. Animal experiments
During the experimental periods, the rats were carefully super-vised for complications including pain, discomfort, signs of infectionand changes in body weight. All animals remained in good healththroughout the study. At the end of the experiment, the rate of animalsurvival was 100%. The animals recovered well after surgery withoutsigns of stress at the time of harvest. The possible minimal differencesin the vital signs that can be objectively measured and observedamong animals did not reach clinically significant levels. Visual mac-roscopic observation revealed that the implants could be easily locat-ed while all of them remained in situ and anchored monocortically.
3.2.1. Subcutaneous implantsTherewere nodifferences in behavior between the alloys of different
compositions and morphology – porous or solids – in the subcutaneousimplants.
Table 5Mechanical properties of the TiZrNb foams.
Foam Porosity(%)
Young's modulus(GPa)
Compressive plateau stress(MPa)
Nb1.6 64 2 46Nb5.5 70 0.3 11Nb1.6a 70 1.4 32
a Values were obtained from Eqs. (1) and (2).
Histologically, the regeneration process was similar in all groupsat 7 days after surgery.
In all groups, connective tissue could be seen in contact with thematerial, with marked vascular proliferation in some areas. Therewas a scarce inflammatory infiltrate. Particularly in Zr–Ti–Nb foamssome macrophages could be depicted, probably due to the porous na-ture of the materials. In this regard, the imprint left by these materialshad more irregular walls due to the intrusion of fibrous tissue in thepores.
In this study, the subcutaneous implantation test was performedin order to assess the cellular response to the materials implanted.Of particular interest was the absence of inflammatory infiltrate andits composition.
3.2.2. Tibial implants
3.2.2.1. 7 days. Decalcified sections were performed in order to ob-serve the cellular response of the host bone against the different ma-terials, particularly the presence or absence of inflammatory reactionand its composition.
Due to the porous nature of the Zr–Ti–Nb foams, after decalcifica-tion, when removing the implants, the scar tissue around the mate-rials was so firmly attached that remained with them, preventingthe observation of the slices of the Zr–Ti–Nb foams. While in theZr–Ti–Nb as remelted this problem did not affect the samples. The in-flammatory response was absent in both groups and a peripheral
327A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
reticular bone response to implants could be observed in the Zr–Ti–Nb as remelted. But in the Nb1.6 as remelted the immature wovenbone was not in contact with the imprint of the material, it was at-tached to a connective interface instead; while immature bone wasin close contact with the imprint left by the implant of the Nb5.5 asremelted.
This observation was corroborated in the undecalcified sliceswhere surrounding the Nb1.6 as remelted, in tibial medullar area, os-teogenic activity with formation of immature bone could be depictedbut not in close contact with metal. On the other hand, although thebone response was similar, it was in close contact with the Nb5.5 asremelted.
With porous implants, in the Nb1.6 foam osteogenesis was scarcewhile in the Nb5.5 foam bone formation was observed partly fillingthe cavities of the material.
The observations made by SEM confirmed the optical microscopydescriptions. Of interest, there were strong differences in the im-plants retrieved from decalcified tissues. While the Zr–Ti–Nb asremelted showed a discrete adherence of the fibrous tissues to thematerial, in the Zr–Ti–Nb foams the tissue almost hid the material,growing inside the cavities (Fig. 9).
3.2.2.2. 60 days. At this time, all samples were undecalcified. In theNb1.6 as remelted it was observed mature bone in contact with thematerial in some areas, while in others the laminar bone was separat-ed from the Nb1.6 as remelted by a fibrous band (Fig. 10a and b),while in the Nb5.5 as remelted mature bone always was observedin close contact with the metal (Fig. 10c and d).
In porous materials mature bone could be observed occupying thecavities but in different proportions, as in the Nb1.6 foam had thelowest amount of bone formation compared to that observed in the
Fig. 9. a) Secondary electrons image showing implant retrieved from decalcified sample atsurface of the material. b) Backscattered electrons image of sample showed in a). Note thaporous material is almost hidden by the fibrous material. d) Backscattered electrons image
Nb5.5 foam, where almost all cavities were occupied by maturebone (Fig. 10e and f). The bony apposition revealed a laminar struc-ture containing individual osteocytes and haversian canals. By SEMsimilar findings were observed.
The rat tibia model was used with the main objective to microscop-ically analyze the behavior of the bone tissue to the tested materials.This model was based on a previously described one for studying theperformance of different types of biomaterials [29] allowing to mimica clinical situation in which an implant material was inserted into thebone. Similar success rates were obtained in the human implant situa-tion, and the rat tibia established the reliability of this method for theplacement of implants [30]. The advantage of this model is the possibil-ity of investigating bone formation in the gap and bone ingrowth in theimplant under very controlled circumstances. The bone healing is notaffected in a purely osseous environment. A disadvantage is the lackof similarity to the clinical situation where such controlled situationsare not found [31].
The absence of a significant inflammatory response in the studygroups confirms the biocompatibility of the materials [32,33]. Follow-ing Borges et al. [34] this is corroborated by the similarity betweengroups in clinical parameters such as dehiscence, infection, pain,lameness, and edema.
Implant placement in bone is presently associated with definedexpectations of success which has been correlated to the histologicalrepresented bone–implant interface and is commonly referred to as‘osseointegration’ [35].
Light microscopic observation showed that all implants have an un-eventful healingwith a normal pattern of bone repair andwithout signsof immunologic and inflammatory reactions. In all instances, directbone-to-implant contact could be achieved during the observation pe-riods. However, the behavior of thematerials was surprisingly different,
7 days after surgery. Note the fibrous tissue which was firmly attached to the poroust almost all implant was covered by the host tissue. c) Higher magnification of a). Theof sample showed in c). Implants are pointed out with arrows.
Fig. 10. Implants at 60 days after surgery. a) The implant Nb1.6 as remelted has bone inclose contact in some areas and in others there is a fibrous interface denoted better bypolarized light (b). c) Mature bone in close contact with Nb5.5 as remelted. d) Same asC observed under polarized light. e) Nb1.6 foam implant with mature bone inside theporous. f) Nb5.5 foam implant, notice the cavities almost filled by lamellar bone. Im-plants are pointed out with arrows. Bones are pointed out with arrows too.
328 A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
independently of their porous or solid nature. Indeed, histological anal-ysis of the bone/implant interface revealed an intimate contact betweenthe Zr–Ti–Nb foams surfaces and the bony implantation bed in thesegroups allowing for a progressive integration of the materials into thebone. At 60 days the bone observed in contact with thematerial was la-mellar or mature. In accordance with our results, other authors [36,37]demonstrated that immature woven bone first appeared around titani-um implants andwas replaced bymature lamellar bone by 56 days afterimplantation in rats, suggesting that osseointegration was almost com-plete at this time.
As expected, in porous materials there was observed a more pro-nounced response than in solid one, in accordance with cell culture re-sults of other authors indicating that the Nb–Ti–Zr foam is morefavorable for cell adhesion and proliferation than its solid counterpart[20].
As stated by Davies and Baldan [38] the regeneration process wassimilar between groups, in subcutaneous as well as, and particularly,
in bone tissue, which also confirms its biocompatibility. However, adifferent behavior was observed in this titanium implants with bio-compatible alloying elements, that is, niobium and zirconium, at dif-ferent concentrations. In this regard, further studies are needed toclarify which element is responsible for the pronounced osteoinduc-tive effects of the Nb5.5 as remelted and the Nb5.5 foam.
4. Conclusions
Ti foams alloyed with non-toxic elements (Zr and Nb) were fabri-cated by means of the space holder method. The following conclu-sions are summarized.
• A high porosity (between 60 and 70%) and an interconnected porousstructure were obtained.
• The average size of macropores was 260 μm.• Young's modulus and compressive plateau stress not only dependon the porosity but also on the amount of Nb in the foam. Young'smodulus between 0.3 and 1.4 GPa and compressive plateau stressbetween 11 and 32 MPa were obtained.
• Nb1.6 foam exhibited a ductile fracture surface morphology whileNb5.5 tended to exhibit more a brittle fracture surface morphology.
• All alloys tested showed excellent biocompatibility in subcutaneousas well as in bone tissues
• Independently of the characteristics of the Nb5.5 alloy (remelted orfoam) it showed pronounced osteoinductive properties.
It can be concluded that it is possible to replace bones with a de-termined porosity with a Zr–Ti–Nb porous alloys because they arefabricated with elements that do not produce adverse tissue reactionand mechanical requirements can be accomplished by changing theamount of the Nb, the minority alloying element.
Acknowledgments
This work has been done under the PICT 2005 No. 32795 contractwith the ANPCYT. The authors also would like to thank Gonzalo Portafor the mechanical tests.
References
[1] G.E. Ryan, A.S. Pandit, D.P. Apatsidis, Biomaterials 29 (2008) 3625–3635.[2] M. Niinomi, Sci. Technol. Adv. Mater. 4 (2003) 445–454.[3] B. Levine, Adv. Eng. Mater. 10 (2008) 788–792.[4] C.E. Wen, M. Mabuchi, Y. Yamada, K. Shimojima, Y. Chino, T. Asahina, Scr. Mater.
45 (2001) 1147–1153.[5] C.E. Wen, Y. Yamada, P.D. Hodgson, Mater. Sci. Eng., C 26 (2006) 1439–1444.[6] J. Banhart, Prog. Mater. Sci. 46 (2001) 559–632.[7] F.H. Froes, Adv. Powder. Metall. Part. Mater. (2001) 84–104.[8] H.C. Hsu, S.C. Wu, Y.C. Sung, W.F. So, J. Alloys Compd. 488 (2009) 279–283.[9] M.V. Oliveira, L.C. Pereira, C.A.A. Cairo, Mater. Res. 5 (2002) 269–273.
[10] Z. Esen, S. Bor, Scr. Mater. 56 (2007) 341–344.[11] E.D. Spoerke, N.G. Murray, H. Li, L.C. Brinson, D.C. Dunand, S.I. Stupp, Acta Biomater.
36–46.[13] Y. Zhentao, Z. Lian, Mater. Sci. Eng. A 438–440 (2006) 391–394.[14] A. Inoue, Acta Mater. 48 (2000) 279–306.[15] M.B. Guglielmotti, S. Renou, R.I. Cabrini, Int. J. Oral Maxillofac. Implants 14 (1999)
565–570.[16] S. Schneider, S.G. Schneider, H. Marques da Silva, C. de Moura Neto, Mater. Res.
8 (2005) 435–438.[17] K. Wang, Mater. Sci. Eng., A 213 (1996) 134–137.[18] M. Long, H.J. Rack, Biomaterials 19 (1998) 1621–1639.[19] M. Abdel-Hady, H. Fuwa, K. Hinoshita, H. Kimura, Y. Shinzato, M. Morinaga, Scr.
Mater. 57 (2007) 1000–1003.[20] X. Wang, Y. Li, P.D. Hodgson, C. Wen, Tissue Eng. Part A 16 (2010) 309–316.[21] C. Schmidt, A.A. Ignatius, L.E. Claes, J. Biomed. Mater. Res. 54 (2001) 209–215.[22] A.I. Pearce, R.G. Richards, S. Milz, E. Schneider, S.G. Pearce, Eur Cell. Mater. 13
(2007) 1–10.[23] M.M. Barreiro, D.R. Grana, G.A. Kokubo, M.I. Luppo, S. Mintzer, G. Vigna, Biomed.
Mater. 5 (2010) 025010.[24] Canadian Council on Animal Care 1993 Guide to the Care and Use of Experimental
329A.E.A. Maya et al. / Materials Science and Engineering C 32 (2012) 321–329
[25] L.J. Gibson, M.F. Ashby, Cellular Solids: Structure and Properties, second ed.Cambridge University Press, Cambridge, UK, 1997.
[26] X. Wang, Y. Li, J. Xiong, P.D. Hodgson, C.E. Wen, Acta Biomater. 5 (2009)3616–3624.
[27] G. Kotan, A.S. Bor, Turk. J. Eng. Environ. Sci. 31 (2007) 149–156.[28] B. Levine, Adv. Eng. Mater. 9 (2008) 788–792.[29] O. Zmener, G. Banegas, C.H. Pameijer, J. Endod. 31 (2005) 457–459.[30] C.M.L. Clokie, H. Warshawsky, Int. J. Oral Maxillofac. Implants 10 (1995) 155–165.[31] D.R. Grana, H.J. Marcos, G.A. Kokubu, Acta Odontol. Latinoam. 21 (2008) 77–83.[32] B. Klinge, P. Alberius, S. Isaksson, J. Jönsson, J. Oral Maxillofac. Surg. 50 (1992)
241–249.
[33] W.R. Walsh, P.J. Chapman-Sheath, S. Cain, J. Debes, W.J. Bruce, M.J. Svehla, E.M.Gillies, J. Orthop. Res. 21 (2003) 655–661.
[35] T. Masuda, P.K. Yliheikkila, D.A. Felton, L.F. Cooper, Int. J. Oral Maxillofac. Implants13 (1998) 17–29.
[36] T. Shirota, K. Donath, Y. Matsui, K. Ohno, K. Michi, J. Oral Implantol. 20 (1994)307–314.
[37] M. Motohashi, T. Shirota, Y. Tokugawa, K. Ohno, K. Michi, A. Yamaguchi, Oral.Surg. Oral. Med. Oral. Pathol. Oral. Radiol. Endod. 87 (1999) 145–152.
[38] J.E. Davies, N. Baldan, J. Biomed. Mater. Res. 36 (1997) 429–440.
![Correlation between microstructure and temperature dependence of magnetic properties in Fe[sub 60]Co[sub 18](Nb,Zr)[sub 6]B[sub 15]Cu[sub 1] alloy series](https://static.documents.pub/doc/80x56/63546f848db8416a940e2684/correlation-between-microstructure-and-temperature-dependence-of-magnetic-properties.jpg)
