SIXTH WORLD CONFERENCE ON TITANIUM, FRANCE 1988

TITANIUM AND TITANIUM ALLOYS, BIOMATERIALS OF PREFERENCE

J. BREME

Lehrstuhl,werkstoffwissenschaft (Mefalle) Universit~t Erlangen-Nurnberg, F.R.G.

1. Introduction

57

Metallic materials which are used as biomaterials have to fulfill the following requirements:

a) corrosion resistance b) biocompatibility c) bioadhesion (bone ingrowth} d) favourable mechanical properties (e.g. Youngs Modulus similar to the bone,

fatigue strength according to the application) e) processability (casting, deformation, powder metallurgy, machinability,

welding, brazing) f) availability (low prices)

Because of these requirements for the application as biomaterials the great num­ber of the metallic materials is limited. Up till now the following groups of metallic materials are. used as biomaterials:

a) stainless steels (e.g. ISO 5832/1 or 316 L) · b) CoCr-alloys (cast ISO 5832/4 or wrought) ISO 5832/6, "Vitallium" c) cp-titanium, ISO 5832/2 titanium alloys, e.g. ISO 5832/3 TiA16V4 d) cp-niobium ' e) cp-tantalum, ASTM 560-78

It is the aim of this contribution to compare the different metallic biomateri­als in use by assis_tance of the_ different requirements.

2. Corrosion resistance In the body fluid, a solution of about 0,9 % NaCl, the pH value amounts to about 7,4 under normal conditions. Changes caused by surgery can effect a rise to 7,8 followed by a drop to a pH of 5,5. After a time of a few days the normal body value of 7,4 is achieved again (1). The most corrosion resistant materials are titanium and alloys, niobium and tantalum followed by wrought Vitallium, cast Vitallium and stainless steel (2, 3). Especially in the passive condition in the body fluid the corrosion current for these metallic material~ is very low, so that only some few ,,ug of metal react per day and per implant, but special con­ditions such as fretting may enhance the corrosion by orders of magnitude. Be­side fretting other corrosion types like crevice corrosion, galvanic corrosion, pitting corrosion, stress or fatigue cracking corrosion may take place in the saline body fluid. In crevi~es e.g. in a screw/plate system the solution pH can drastically be decreased towards values of pH= 1. Fig. 1 shows the current density of different materials as a function of the po­tential difference between the anodic and cathodic branches of the current po­tential curves in 0,9 % NaCl with a stable redox system Fe(CN)64-/Fe(CN)63- (2). The saline containing this redox system resembled closely in its resting po­tential that of a tissue culture fluid, which has its redox potential at 400 mV. Ti and alloys, Ta and Nb behave .more noble ·than the stainless steel AISI 3161 and a wrought CoNiCr-alloy. The same range can be observed during the measure­ment of the polarization resistance of the different materials (Table 1) (2). Break down potential measurements of different implant materials in Hank's solu­tion resulted also in a clear order of rankirig of the different materials. While commercially pure titanium and TiA16V4 had high breakdown potentials of 2,4 and 2,0 V respectively, for stainless steel and CoCr-alloys (cast and wrought) this value amounted only to 0,2 and 0,42 V respectively (Table 2) (3). As already described in a former publication (4) Ti and its alloys, Nb and Ta belong to the group of metals which in body fluids cannot undergo a breakdown of passity. In this fluid a breakdown at a high potential causing pitting corrosion is impos­sible, because it is more positive than the oxygen reduction reversible poten­tial. On the other hand the passivation potential.is less positive than the wa-

58

.tJ.f(mVJ---

Fig. 1: Current density as a function of the potential difference between the anodic and cathodic branches of the current-potential curves for metals tested in 0,9 % NaCl with a stable re-dox system [Fe(CN)64-/Fe(CN)63-) ( 2)

ter or hydrogen-ion reduction.

Table 1: Polarisation resistance of me­tallic biomaterials in 0,9 % NaCl with a stable redox system [Fe(CN)64-/ Fe(CN)63-J ( 2) '

R (kncm2 ) p Au 0,28

FeCrNiMo ( 316L) 4,38

CoNiCr (wrought) 3,32

cp-Ti 714

TiA16V4 455

cp-Nb 455

cp-Ta 1430

In all materials the passive layer can be damaged mechanically e.g. by fretting metal on metal (plate/screw) or by the instruments used during surgery. The time of the repassivation of the material is the·refore very important. The re­passivation behaviour of different materials in saline solution.was measured using an electrode which rotates by 10 sec-1 in saline solution whereby it is activated by a cutting tool of Al203. The decrease of the corrosion current is measured in dependence on the time at different potentials. The repassivation

Table 2: Breakdown potential in Hank's solution of metallic biomaterials (3) and repassivation time in 0;9 %.NaCl

breakdown potential (V) repassivation time (msec) (calomel electrode) ( 3) te to 05

-0,5 v +O, 5 V -0,5 V ' +O, 5 FeCrNiMo ( 316L) +O, 2 .- 0,3 >72000 35 »72000 .>6000 Co Cr (cast) +0,42 44,4 .36. »6000. >6000 CoNiCr (wrought) +0,42 35,5' 41 >6000 ·5300 TiA16V4 +2,0 37 41 43,3 45,8 cp-Ti +2,4 43 44,4 47,4 49 cp-Ta +2,25 cp-Nb 47,6 43, 1 47 85

v

is defined to be achieved, if the current. density will. amount to 1/e (eni2,718) of the current density in the activated condition (5). In additiOn the time to 05, of a rest active current density of 5 % was determined. The values te and to;o5 of .the different measured materials are given in Table 2. The passive oxygen surface layer (te) is reconstructed dependent on the material in some milliseconds. The growth of the surface layer (to 05) of cp-titanium and the ti­tanium alloys is accelerated compared with the other materials. In order to avoid a damage of the surface layer.coatings with hard layers of not abrasive materials which have in addition a favourable fretting behaviour, are recommended, especially for movable parts of implants. Since highest values of the acceleration tensi6n are achieved by ion implantation, the best .reaction and binding can:be expected with this procedure. By ion implantation of TiN on

59

wrought Vitallium e.s. beside the fretting behaviour the corrosion resistance of the material tested in 0,17 M saline solution was improved. The pitting of the surface treated material amounted to 1,16 V while the material without a surface treatment had a potential of 0,83 V (6). If there are cracks and fis­sures in the surface layer the corrosion rate will be accelerated due to the lower pH value in these crevices. A principle question which arises, is the ne­cessity of a surface treatment of materials by TiN, since implants of Ti and its alloys are available which themself can easier and with more security be treated by nitr6gen and which in addition offer the best co~rosion resistance. Experi­ments with ion implantation of nitrogen in titanium surfaces had a good result concerning the fretting behaviour. Even the fatigue streneth of the alloy TiA16V4, which was surface treated by nitrogen ion ifuplantation, is reported to be increased due to compression stresses generated by the high acceleration ten­sion of the nitrogen ions (7). Another possibility to harden the surface of Ti and its alloys without diminish­ing the corrosion behaviour and fatigue properties is a short annealing in air e.g. by an induction heating and subsequent quenching. This method was applica­ted for the improvement of the friction behaviour of heads of hip prosthesis ( 8) .

3. Biocompatibility In the system implant/body several interactions generating injuries are possi­ble:

a) by the corrosion process a flow of electrones in the implant metal and a flow of ions in the surrounding tissue is produced. This ion flow in the tissue may disturb the physiological ion movement of the nerve cells

b) inorganic reaction of the implant or of primary corrosion products by the solution of metal ions in the body fluid and transport to the different organs where they are concentrated and can produce systemic or hypersensi­tive effects, if the limit of toxicity for a certain metal will be excee­ded

c) organic, direct reaction of the implant or of primary corrosion products with proteins of the tissue ccmsing e.g. an inflammation

d) generation of H202 by inflammatory cells and decomposition of H202 by for-matirJ of a hydroxyl radical causes injury in the biological system.

Whether one of these interactions occurs or not, depends on the physical and chemical properties of the different materials. Ti, Ta and Nb are reported to be biocompatible, because they are forming protective surface layers of semi- or nonconductive oxides. These oxides are able to prevent extensively an exchange of elec~rons and therefore a flow of ions through the tissue (9) due to their isolating effect. This isolating effect is demonstrated by the dielectric con­stants of the different metal oxides (Table 3). There are three groups of ox­ides. While Ti02(rutile), Fe203 and Nb205 have constants even higher than that of water, Al203, Cr203 and Ta205 have a lower isolating effect and a better con-. ductivity (10). For Ni and V-oxides dielectric constants are not available be­cause of their high conductivity. The relatively low isolating effect of Ta­oxide is indirectly proved by cytotronic effects of Ta on the membrane proper­ties and on the growth of spinal ganglion cells during in vitro tests. In con­trast Ti showed no effect on the membrane properties and on the growth of · ganglion cells (11). Due to the isolating properties of Ti-oxide implants of Ti are not recognized by the bone or tissue as foreign bodies. Concerning inorganic or organic reactions the primary corrosion products of the metallic implants are mainly responsible for the biocompatibility of the implanted metal, because they can have due to their great ~urface favourably an interaction with the tissue or with the body fluid. By a solution in the body fluid the metal is transported to the different organs, where by an enrichment of these metals an undesired interaction can arise. The primary corrosion products of the most important elements in metallic implant materials have a different thermodynamic stability. While the oxides or hydroxides of Al, Cr, Nb, Ta, Ti and V are stable due to a more negative heat of formation than that of water,the oxides and hyroxides of Co and Ni are unstable because of a less negative heat of formation than that of water (Table 3) (12). The interaction between the oxide or hydroxide and the body fluid is increased, if the heat of formation for the oxide or hydroxide is increased. Therefore the thermodynamically stable corrosion products have a low solution product and a low solubility in the body fluid. This is directly demonstrated by the pk-values (negative logarithm) of the solution product 'Of the primary corrosion products (Table 3). While Ti-, Ta-, Nb- and Cr-oxides have pk~values >14 i.e. hydrolisis

60

Table 3: Dielectric constant E,heat of formation 6H and solubility Pk (neg. Log.) of primary corrosion products compared with H2o

primary 0

corrosion E ( 10) -t.H298

pk( 1 3) product KJ/mol

( 1 2)

Al203 5-10 1675 +14,6

Al(OH)3 916

coo 239 -12,6

Cr203 1 2 1141 +18,6

Cr03 595

Cr(OH)3 988 - 1'8

FeO 267 -13,3

Fe203 100 822 -14

Fe(OH)2 30-38 568 + 2,3

Mo03 712 + 3,7

NiO '.'40 -12,2 Ni(OH)2 538

NbO 486 Nb205 280 1905 >20

Ta205 1 2 2090 >20

TiO 518 Ti02 anat. 48 935

brook 78 +18 ruti 11 110 943

VO 410 V205 1560 +10,3

H20 78 273 +14

lo 60

[r 40 50

~ Ni l,f)

30 15

,0

lO 10

20

10 10

;5 15 15

• inurmelµglgcrei1tmme) •mblood(nglgJ omplasmafnglgJ months

Fig. 2: Metal clearance in a patient with a total hip prosthesis of a CoCr-alloy after removal (The horizontal lines indicate the upper normal values) (16).

w ti u z ;'!: 7

SEOUESlRATION

3161 ~ • Vl Vl w

Co Cr MoN1 TOXICllY

er 6

z • ~ v c::> s ~ ' >- , [ Cc' r" <{ f

1

~ B e Al

~ : ~--~-c_o _____ c_FQ_M_o ___ _

O> 0 ;, SSvE REACTION

T1 ALLOYS

Fig. 3: Polarization restances grouped ac­cording to tissue reaction of metals (13).

cannot play a role, Co-, Fe-, and Ni-oxides posses even negative Pk-values which cause a considerable solubility. In spite of a high negative heat of formation for Fe203 and Fe- and Cr-hydroxide negative pk-values and a high solubility are reported (13). A remarkable solubility of Cr in serum was observed (14) while titanium is due to the formation of the thermodynamically very stable 'I'i02 practically unsoluble. Measuring the concentration of different metals in vari­ous organs six and sixteen weeks after the implantation the content of titanium amounted after 6 weeks to 451 ppm in the spleen and to 53,4 ppm in the lung of a rabbit. After 16 weeks a decrease to 13 ppm and 8 ppm respectively was measured, These values correspond to the values in the normal spleen and lung. No signi­ficant changes in the liver and in the kidneys were observed. In contrast Co and Ni from cobalt-based c.lloys and stainless steel were found in higher concen­tration in-these organs (15). Patients with total hip replacements by implants of stainless steel or CoCr-alloys who got difficulties after 2 to 15 years by loosening of the prosthesis and/or allergical reactions to Cr, Co or, Ni had an increased content of these elements in urin, plasma and blood. Even 15 months after removal the contents of these elements was excessive in urin, blood and plasma (Fig. 2) (16). Table 4 gives the level of toxicity of different elements in the kidneys. This

Table 4: Limit of toxicity CCR50 of different metal salts (17).

v Cr Ni Co Mn Fe

CCR50 (µg/ml) 1 ' 1 3,5 15 59

level was determined by investigating the reaction of salts of different ele­ments with cells of the kidney of green African monkies. The so called CCR50-

61

value was measured which is defined as the concentration of the studied sub­stance which generates a reduction of survival of the renal

2cells of 5.0. %. For

vanadium the lowest value of all measured elements of 3·10- µg/ml was observed (17). This is the reason why absolutely biocompatible biomaterials which do not contain toxic elements such as the alloy TiA15Fe2,5 (18) and TiA15Nb7 (19) were developed. Beside the inorganic reaction of the metal ions other reactions with organic constituents in the tissue with proteins are possible. Three different reactions may happen:

a) the coagulation and precipitation of proteins which is produced by an electrostatic effect of the metal ions. Therefor~ Cr3+ and Fe3+ ions have a greater influence than cr2+ and Fe2+.

b) the formation of complex bindings between the proteins and the metal ions according to the equation

( 1 ) ChH2 + Me2+ - ChMe + 2H+ whereby Ch is the protein which forms a chelat

c) the ~ncapsulation of solid corrosion products by proteins. There are two reactions in concurrency which can be estimated by the equation

(2) CheM + 2H20 - Me(OH)2 + ChH2 If the precipitation of hydroxides is stronger than the formation of complexes, the equation acts to the right side. On the other hand the precipitation of hy­droxides can be avoided by the formation of complexes. A strong complex forming protein should also be able to dissolve an already precipitated hydroxide (14). Thermodynamically stable primary corrosion products with a low solubility in the body fluid are in a stable equilibrium with only a low reactivity against the proteins of the surrounding tissue. With materials which behave inert or biocompatible the cells in the vicinity of the implant were still supplied with blood, while cells in the neighbourhood of toxic materials showed an inflamma­tory reaction and died off, Fig. 3 shows the result of this classification and the values of the polarisation resistance of the different materials. Some few elements (Cr, Co, Ni and V) show toxic effects and are also connected with a relatively.low polarisation resistance. Ti and-its alloys, Nb and Ta which have a high·polarisation resistance, behave inert. Inbetween the materials appear which are capsulated. The result of Fig. 3 shows also, that the corrosion be­haviour given by the polarization resistance is not only responsible for the biocompatibility of the material exposed to the tissue. The 3161-steel and the CoCr-alloy which have a similar polarization resistance to titanium, are capsu­lated by a tissue membrane and do not behave inert (13). With the insertion of an implant an inflammatory reaction is associated. H202 is generated by inflam­matory cells (20, 21). By the reaction of hydrogen peroxide with metal ions a hydroxyl radical (OH')(is formed according to

(3) en+ + H2o2 - Me n+1 >.+ + OH- + OH' This radical is able to cause injury in biological systems e.g. biomembranes can be deteriorated. Titanium is able to bind H202 in a Ti-H202 complex. This complex can trap the superoxid radical which is formed·during the H202 decompo­sition. By spectrophotometric spin-trapping.measurements and electron spin re­sonance measurements no hydroxyl radical formation rate in Ti-H202 could be de­tected. A similar result was observed with Zr, Au and Al (22).

4. Bioadhesion (bone ingrowth) _ The integration of metallic implants by an ingrowth was studied for a lot of different materials and implant systems. The ingrowth behaviour of miniplates of commercially pure titanium and of the stainless steel 3161 was investigated by the implantation of these plates to the legs of Hanford minipigs. The mini­plates were fixed to the legs of these pigs by screws. After the removal after an exposure time of 8 weeks a hi~tologic examination by fluorescence microsco­py was performed. In all animals where titanium plates were used a new bone formation could be observed in close contact to the surface of the screws and plates (Fig. 4). In contrast to this result the new bone formation using stainless steel was less and in addition a granulated tissue between the metal­lic surface and·the surrounding bone was found (Fig. 5) (23). 'rhis granulated tissue at the interface bone/implant has the disadvantage, that it is not sup­plied by blood. Therefore a systematic treatment of the host tissue against in­flammatory reactions in the vicinity of the implant by injections Vlill have no success, because the antidotes cannot be transported directly to the place of the inflammation. In addition the granulated connective tissue is not able to transfer or sustain forces, so that a loosening of the implant will take place. A growth of the bone in close contact was already reported in a lot of former

62

0.35mm 1----t -~

Fig.4: Histologic examina­tion by fluorescence mi­croscopy of a Ti-plate, 8 weeks after implantation to the bone of a Hanford pig (23)

Fig. 5: Histologic exami­nation by fluorescence microscopy of a 3161-plate 8 weeks after im­plantation to the bone of a Hanford pig (23)

I I . ' t '·· i·' .. :. rt

Fig. 6: Cell reaction with beads of TiMo10 (31)

investigations (24-29), which studied in detail the contact area tissue/implant. From all types of implants the dental implants have to fulfill the most critical requirements, because they are in contact with three different types of tissue, Beside an attachment to the alveolar bone (hard tissue) and to the periodontal ligament (connective tissue) for bone maintenance an attachment to the gingiva (epithelium) for seal is the key problem. This sealing of the part of the im­plant which enters the oral cavity, has to prevent inflammatory reactions caused by bacteria. DUe to the good ingrowth behaviour implants of titanium show no tendency of the downgrowth of the epithelium which provides inflammatory reac­tions in the interfacial zone (30), The adhesion of titanium materials to the soft tissue was proved by culturing Hela cells and Hp (pulp) cells with beads of titanium and its alloys in a gyratory shaking incubator. After 24 hours culture cells adhered already to the surface of this beads. After 7 days the titanium beads were closely bound with cell to alloy and cell to cell adhesion. The cells showed cytoplasmic adhesion and cytoplasmic bridges were formed among the beads (Fig. 6) (31). In contrast using stainless steel as implant material a fibrous encapsulation which separates the implant from the surrounding tissue is formed. This fibrous encapsulation replaces the intercellular ground substance observed with titanium implants. A similar unfavourable behaviour was found with dental implants of a CoCr-alloy in dogs. Histological findings showed that 28 days af­ter the implantation new formed bone fibrils did grow to the surface of the me­tal. But already after 56 days by a decrease of the new formed bone caused by a resorption the hole of the implant enlarged and after 112 days the implant of Vitallium got lost. In contrast an implant of Vitallium which was plasma coated with titanium, again had a close contact to the bone (32). In order to ensure a perfect integration the titanium implant must be unloaded during a period of about 3-4 weeks. An initial implant movement relative to the host bone can re­sult in an attachment by a nonmineralized fibrous connective tissue layer with a poor adhesion strength. The ranges of movement which resulted in either bone or fibrous connective tissue fixation, were observed with dental implants in dogs. An osseointegration was observed, if the movement of the implant relative to the bone did not exceed 28 )4m, while an excess movement of 150 p.m or more caused an attachment by connective fibrous tissue (33). Using two dental im­plants in animals whereby one sustained immediately functional loads while the other one remained nonfunctional it was demonstrated that even minute movements can cause the induction of a fibrous capsule. An implant initially surrounded by such a fibrous capsule cannot be expected to become properly osseointegrated if the loading conditions are later optimized (31). After the ingrowth period a loading of the implant is desired in order to transmit the pressure whereby a

63

new bone formation is stimulated. Fig. 7 shows the results of the adhesion strength of titanium implants to the ulna of monkeys. ·with increasing time after implantation the strength increased and achieved its maximum value after 16 to 17 weeks. In addition an influence of the surface structure was found. A titanium sand blast surface· has a better ad­hesion strength compared with plasma titanium coated surface (34). In another study the influence of the surface roughness of cylinders of titanium and tita­nium alloys, which were implanted to the legs of rabbits, .. was investigated. Fig .. 8 shows the ~esults of bioglass in comparison with TiA16V4 and TiA15Fe2,5. A measurable adhesion of the titanium alloys could only be observed, if acer-

c: ~ 2 QI

.i:::. u 0

o Ti sand blast • Ti plasma coated 0

100 200 time after implantation [ d ]

Fig.7: Adhesion strength of Ti-plates to the ulna of monkies after different time of expo­sure ( 34).

Tensile strength of the rnterf~ce bone I implant

'E I Sch~itz H J . Gross U et al 1988 I .§ z

TiAl5Fe2.5 l168 d) , I £ 1.8

0

//'

0

TiAl6V4 1168 di a> c ~ ;;; 1.6

.'!! / iii

1.4 I c 2 I

1.2 I I TrAl5Fe2.5 184 di

I ii I I x

I 1.0 I I "-.. I bioglass 184 d I I I a.a I I

06 I I

....-! 1/ ,.,. TiAJ6V4 184 d I ,.,.

/I ,.,.

0.4 ,.,. I/ ,,,. ,.,.

/ 0.2 It /'' J time of exposure in days .

/

. 10 20 30 40 50 implant surface roughness lµm I

Fig.8: Influence of the surface roughness of implants on the ten­sile strength of the interface bone/implant of different ma­terials (35)

tain surface roughness ( >22 p.m) of the implant was given. With increasing roughness the adhesion strength is improved. The time after the implantation has also an influence. With implants of TiA16V4 the tension strength for tearing off the cylinders from the bone was more than doubled, if the time was increased from 84 days to 168 days. After the short exposure time of 84 days the implant of TiA15Fe2,5 had already an adhesion to the bone similar to an implant of bio­glass. In contrast to the titanium alloys the adhesion of the bioglass was not dependent on the surface roughness (35). These results show, that the growth of the bone and tissue in close contact to Ti and its alloys under the formation of a strong bond must have more a biomechanical than a chemical bioactive charac­ter. Consequently it was shown, that by an increase of the surface area e.g. by drilling holes (Fig. 9 and Fig. 10) in the contact area of the implanted cylin­ders to the bone the tear off force necessary was increased. But taking in con­sideration the supplementary surface the strength of adhesion was not increased. A hydroxy-apatite coated TiA15Fe2,5 implant showed already 84 days after ~he im­plantation a maximum adhesion strength of 1,97 N/mm2 (35). The better fixation of the bone at a structural implant surface leads conse­quently to a porous implant, whic~ allows an ingrowth of the bone. Beside a bet­ter fixation there are two additional reasons for the porous implant: Youngs Mo­dulus of the implant is decreased whereby the functional load is better trans­mitted and a new bone formation is more stimulated. In addition the damping ca­pacity of the implant is increased and the shearstress generated by the func­tional loading of the implant at the interface implant/bone is decreased, be­cause similar to the thread of a screw the load at this interface causes a nor­mal stress perpendicular to the inclined area and a smaller shear stress which is effective in the inclined area. Whether the ingrowth of the tissue in a po­rous implant remains fibrous or becomes mineralized depends on the pore size. As

64

shown with porous ceramic implants below a pore size of 100.,.u.m soft tissue is allowed to grow in, whereas pores of a larger diameter allow the development of mineralized bone (36). A comparison of a surface coated titanium (40 - 50 % po­rosity, pore size 75 ,µm) implant with a smooth implant of titanium embedded in the femura of sheeps showed the superiority of the porous coatea implant in its

Fi~. 9: Cylinders with ho es for the implanta­tion to the legs of rabbits (35)

Fig. 10: Bone after measuring the adhesion strength to cylinders of Fig. 9 (35)

40

30

Adhesion strength to epithelial cells

I Fletcher et al 1979 I

20 .;;i-----0----------.0 Cu,A9 o--

2 3 time of exposure l h ]

5

Fig. 11: Adhesion strength of epithelial cells to different materials, influence of the time of exposure (38).

bond strength. 26 weeks after the implantation for the implant with the porous surface a shear strength of 17,5 N/mm2 was measured, while the uncoated im~lant had a shear strength of only 0,12 N/mm2 after the same period of exposure (37). The adhesion of gingival cell lines which is demanded for dental implants was tested in vitro with different materials. For this purpose cells were grown on different rigid material samples with a surface of about 1 cm2 which were per­forated with a 0,46 mm hole. The degree of cell adhesion was investigated by measuring the air pressure required to debond the materials from the cells. Fig. 11 shows the results. Titanium, bioglass and Vitallium provided the best mate­rials (in this given order of ranking). Copper and silver showed poor results, probably because of their toxic effect on cells (38). The strong bond of Ti im­plants to the bone and to the soft tissue can have different reasons. The high dielectric constants of titanium oxides give an increased van der Waals bonding. ~he similarity to water may be seen as a rather natural environment by the bio­molecules. Ionic and covalent bonds which can arise from overlapping charge clouds of the surface atoms and the biomolecules, are also suggested with tita­nium oxide. In the oxide layer mineral ions from the biosystem e.g. calcium and phosphorus are incorporated (31,34) by a diffusion via lattice defects. These Ca- and P-deposits are able to bridge the gap between the collagen fibrils and the oxide surface layer with maximum mechanical stability. Indeed with dental implants removed from animals it was possible to demonstrate hydroxyapatite crystals which bridged the distance between the tissue and the implant (40). Therefore after tearing off the implant from the bone on the surface of the im­plant collagen fibers often were observed (41). The test of the adhesion strength of the bone to a cylinder with holes (Fig. 9) demonstrated a bonding of the material to the bone. Only some few bone bars grown in the holes did not fracture during this test (Fig. 10) (35). By these considerations it is obvious, that titanium materials are not only osseointegrated due to their bioinertness, which allows a mechanical bonding between the bone and the implant, but also due to a certain bioimitating bioac­tivity, which has a chemical character. Compared with apatite or bioglasses the mechanical aspect of the osseointegration should be predominant for titanium.

5. Mechanical properties Table 5 shows typical values of the mechanical properties for biomaterials. Be­side a Youngs Modulus similar to the modulus of the bone (10 000 - 30 000 N/

65

mm2) a sufficient fatigue strength and elongation at fracture As are deman-ded. Ti and its alloys and Nb have an elastic modulus of about 100 000 - 110 000 N/mm2. This value is from all metallic biomaterials the closest by the modulus of the bone. In their fatigue strength titanium and its alloys are equal or even superior to other commercially pure materials and alloys respectively. Consid­ering the value of the biofunctionality in Table 5 titanium and its alloys de­monstrate their superiority to other biomaterials. By the "biofunctionality"

Table 5: Mechanical properties (typical values)

X2CrNiMo18 1 2 Co Cr CoNiCr TiA16V4 TiA15Fe2, 5 cp-Ti cp-Nb cp-Ta AISI 316L (cast) (wrought

210 200 220 105 105 100 120 200 Young's Mod. ·103 (N/mm2)

450 500 850 900 900 300 250 300 RpO, 2 lN/mm<:'.)

250 300 500 550 550 200 150 200 ab* (N/mm2 )

40 8 20 13 15 30 70 40 A5 (%)

*rotating bending fatigue strength

the great importance of Youngs modulus is taken into account. The lower Youngs Modulus the better the functional load on the implant can be transmitted, where­by the formatio·n of new bone is sti:::ulated. Allowing a higher elastic deforma­tion of the implant a higher pressure is transferred to the bone. This fact was indirectly proven by demonstrating, that in areas of the implant, where no load was active, in the so called "load shadow lines", only soft tissue could be ge­nerated (42). Since titanium and its alloys offer a wide range of mechanical properties (Table 5), the application in surgery is extensive. While high loa­ded devices are fabricated of titanium alloys, low bearing devices are produced of commercially pure titanium in the as delivered or in the as delivered and cold deformed condition. Typical examples for high loaded devices are hi~ pros­thesis, knee joints, the replacement of the bone of an upperarm (Fig. 12) or of a tigh (Fig. 13). Typical examples for the application of cp-titanium in the

c=•- /.'

Fig. 12: Device (TiA15Fe2,5) for the re- Fig. 13: Device (TiA15Fe2,5) for the placement of a part of the upper arm replacement of a part of the tigh

cold deformed condition are plates and screws, bundle nails (43) and Kuntscher­nails. In oral surgery beside dental implants meshs (Fig. 14) (44) and plate­screw systems of titanium materials are used for maxillofacial reconstruction (Fig. 15) (45). Because of notch effects the fatigue properties of implants with porous surfaces are decreased. From the different production procedures of these surfaces plasma spraying, sintering and diffusion bonding the best fatigue strength which is about 10 % lower than the fatigue strength of the bulk mate­rial was achieved by plasma spraying (46).

66

Fig. 14: Maxillofacial reconstruction, plate-screw system of cp-Ti (45)

6. Processing Melting, casting, deformation, 'processing by powder metallurgy, machining, bra­zing and welding of Ti and its alloys are possible and are used for the produc­tion of the different implants. While the stems of hip prosthesis e.g. are die­forged in standard dimensions or machined individually from semifinished mate­rial by CNC milling tool machines, the sockets of the hip prosthesis can be pro­duced by investment precision casting with a favourable fine grained micro­structure. A particular problem of dental implants is the joining of alumina ceramic with titanium in order to produce a composite material. Such a dental implant which consists of a metal post inserted in the alveolar bone and of a ceramic part passing through the gingiva in the oral cavity, combines the advantages of both materials. It was shown that in the area of the gingiva titanium implants are often affected by plaque which can generate inflammation. In comparison with alumina at titanium implants the number of microorganisms was increased by plaque (48). On the other hand compared with alumina the titanium post has ex­cellent fatigue properties and is able to withstand tension stresses. In order to introduce no incompatible materials the active diffusion welding is chosen as joining procedure. This method is performed under vacuum at a temperature of 1300 °c, whereby under load a direct contact between titanium and alumina is formed. First welding tests with commercially pure titanium, TiA15Fe2,5 and several TiAl and TiMn-alloys showed that a good weldability with alumina can be achieved. But because of the thermal misfit between Ti and alumina during cool­ing cracks occurred in the ceramic produced by tension stresses. By alloying the

o Ti

12 e TiTa10

• TiTa30

10 I /I'',, .. ",

l ~ .... . ir ,\

I .

• 100 200 300 400 500 600 700 600 900 1000

'c

Fig, 16: Thermal expansion of different TiTa- Fig. 17: Diffusion welded Al203 alloys in comparison with cp-Ti and Al 2o3 with 6iTa40, welding temperature

1300 c (50)

67

thermal expansion coefficient of Ti was adapted to the coefficient of alumina. Since only biocompatible elements are suitable, TiTa-alloys with 30-40 % tanta­lum were developed for this purpose (49). Fig. 16 shows the expansion coeffi­cient of TiTa-alloys dependent on the Ta content. With alloys with 30 and 40 % tantalum a welding without cracks with alumina was possible (Fig. 17).

7. Conclusion Table 6 gives a summary of the requirements for metallic biomaterial consider­ing the different materials. Titanium and its alloys must be the biomaterials of choice, because they possess the best corrosion resistance, biocompatibility,

Table 6: Comparison of metallic biomaterials

Corrosion resistance Biocompatibility

Bioadhesion

Processabili ty •• 1 '2 C D w p

1 t 5 c w p

+ + ( +)V?

+ 2,3 5,2 C D C D w p w p

+

+

+ 5,2 C D w p

+

+

+ 1 '8

C D w p

+

+ ( +)?

1 '3 C D w p

+ +

(+)?

1 '3 C D w p

Costs*" ( DM/kg) N60 rv70 N75 ... so rv70 1\1300 N450

• Biofunctionality •• C =Casting; D = Deformation; W = Welding;P =Powder Metallurgyar.e possible ... semifinished product

bioinertness, osseointegration and biofunctionality of all materials. The pro­cessability of titanium and its alloys is unrestricted. Considering the economic aspects (availability) the titanium materials belong to one group with stain­less steel and CoCr-alloys, while tantalum and niobium belong to another more expensive group. Due to these facts titanium and its alloys are the biomaterial of preference. '

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