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Fatigue characteristics of metallic biomaterials
International Journalof Fatigue
International Journal of Fatigue 29 (2007) 992–1000
www.elsevier.com/locate/ijfatigue
Fatigue characteristics of metallic biomaterials M. Niinomi
*
Department of Biomaterials Science, Institute for Materials Research, Tohoku University, 2-1-1, Katahira Aoba-ku, Sendai, Miyagi 980-8577, Japan Received 16 November 2005; received in revised form 30 August 2006; accepted 26 September 2006 Available online 13 November 2006
Abstract It is essential to determine the fatigue characteristics of metallic biomaterials for long-term safe usage of implants in living body such as hip joints, bone plates, and dental implants. On the basis of reported data, the fatigue strength, fretting fatigue strength, and fatigue crack propagation rate of metallic biomaterials – mainly titanium alloys, Co alloys, and stainless steels – in air, a living body and a simulated body environment are described extensively with respect to related microstructures. Further, the fatigue strength of titanium alloys coated with bioactive ceramic and that of biometallic wire are examined. Ó 2006 Elsevier Ltd. All rights reserved. Keywords: Fatigue; Fatigue crack; Fretting fatigue; Metallic biomaterials; Titanium alloy; Co alloy; Stainless steel
1. Introduction Implants that function as bones – for example, implants, which replace failed hard tissue, such as artificial hip joints, artificial joints, bone plates, and dental implants – are usually used under severe cyclic loading conditions. Therefore, metallic materials that typically exhibit high strength, ductility and toughness are the main candidates for the structural biomaterials of these implants. Further, the abovementioned implants must exhibit high biocompatibility, high performance, and reliability for long-term use. The reliability of the implants after implantations is determined by their fracture and wear after the critical period of infection. Therefore, the mechanical properties of structural biomaterials in a living body environment – such as fatigue, toughness, and wear resistance – need to be evaluated and then improved in order to confidently use the implants for a long period of time. With regard to the fracture of structural biomaterials, fatigue fracture occurs occasionally; it is considered to be a crucial problem among the various types of fractures.
*
Tel.: +81 22 215 2574; fax: +81 22 215 2533. E-mail address: [email protected]
0142-1123/$ - see front matter Ó 2006 Elsevier Ltd. All rights reserved. doi:10.1016/j.ijfatigue.2006.09.021
Fatigue with fretting, i.e., fretting fatigue, is a type of fatigue that can occur between two bodies, such as between a bone plate and screw. Fatigue characteristics are closely related with the microstructures. The microstructures in metallic structural biomaterials change according to the processing and heat treatment employed. Therefore, in this paper, the fatigue characteristics of metallic structural biomaterials are described extensively with respect to related microstructures or processing for microstructural control. 2. Fatigue strength in air Fig. 1 [1] shows an example of the fatigue limits of stainless steels, Co alloy, and titanium and its alloys as representative metallic biomaterials in air; the fatigue limits are scattered depending on factors such the fabricating process, surface condition, microstructure, and fatigue condition. The fatigue limit of bovine bone [2] is also shown in the Fig. 1. This limit decreases in the following order: Co alloy P Ti–6Al–4V P 316L stainless steel. However, the limit of each metallic biomaterial shows a fairly large scatter due to the abovementioned factors. The fatigue limit of each metallic biomaterial is higher than that of bovine bone. Figs. 2 and 3 [3,4] show the relationship between the rotating bending fatigue limit and the elongation up to fracture and that between the fatigue strength and Young’s
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
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Fig. 1. Fatigue strength at 107 cycles of biomedical stainless steel, Co alloys, titanium and it’s alloys, and bone. Data without designation of rotating bending are those obtained from uniaxial fatigue tests.
Fig. 2. Elongation at fracture as a function of the fatigue strength of metallic biomaterials.
modulus for various metallic biomaterials, respectively. The rotating bending fatigue strength increases with the elongation up to fracture. The rotating bending fatigue strength is relatively higher and Young’s modulus is relatively low in a + b titanium alloys. Therefore, the BF value of titanium alloys, which is obtained by dividing the fatigue strength with Young’s modulus, is high (.5.2). The BF values of materials such as 316L stainless steel, Co–Cr alloy, Co–Ni–Cr alloy, pure Ti, and pure Nb, lie between 1.2 and 2.3 [3]. This indicates that the mechanical biocompatibility of titanium alloys is excellent in comparison with that of other metallic biomaterials. The BF value of aged Ti–29Nb–13Ta–4.6Zr (TNTZ alloy), which has recently been proposed for use in biomedical applications, is very high (.9.6) [4].
Fig. 3. Fatigue strength and Young’s modulus of each metallic biomaterial. TNTZ means low modulus b type titanium alloy, Ti–29Nb–13Ta– 4.6Zr.
Among the titanium alloys that have received considerable attention as metallic biomaterials, the fatigue limit of pure Ti, which is a type alloy, is the lowest. The fatigue limits of Ti–6Al–4V and Ti–6Al–7Nb, which are a + b type alloys, are among the highest. The fatigue limit of solutionized b type alloy is low; however, it increases drastically to approximate that of an a + b type alloy by aging after solutionizing due to precipitation of the a phase. The fatigue strength of a + b and b type alloys varies considerably with
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
Air Ringer’s, f=10 Hz
Ti-6Al-4V ELI : A (Equiaxed ) : C (Windmanstatten
700
(MPa)
600 500
400
300
200
100 104
3. Fatigue strength in simulated body environment In order to estimate the fatigue strength of metallic biomaterials in a living body environment, it is essential to evaluate it in a simulated body environment. Fig. 5 [6] compares the rotating bending fatigue strength of Ti–6Al–4V ELI and SUS 316L stainless steel in Ringer’s solution with those in air. The fatigue strength of Ti–6Al– 4V ELI in Ringer’s solution is similar to that in air, while that of SUS 316L stainless steel in Ringer’s solution is lower than that in air. Therefore, corrosion fatigue occurs in Ringer’s solution for SUS 316L stainless steel. The fatigue strength of Co–Cr–Mo is also reported to decrease in Ringer’s solution as shown in Fig. 6 [7]. As shown in Fig. 7 [6], the fatigue strength of Ti–5Al– 2.5Fe in Ringer’s solution decreases when the oxygen con-
)
: SUS 316L
max
the microstructure. In a + b type alloys, the fatigue strength of an equiaxed a structure is, in general, higher than that of an acicular or Widmansta¨tten a structure [5]. It is possible to improve the fatigue strength of b type alloys by thermomechanical treatment involving of cold working and heat treatment. Cold working can be performed very easily in b type alloys. Fig. 4 [4] shows the fatigue strength of low modulus Ti–29Nb–13Ta–4.6Zr – which was obtained after solutionizing, aging (TNTZST aged at 598 K and TNTZST aged at 673 K), severe cold rolling and aging (TNTZCR aged at 598 K and TNTZCR aged at 673 K) – with the fatigue limits of Ti–6Al–4V ELI and Ti–6Al–7Nb. The fatigue limit of Ti–29Nb–13Ta–4.6Zr is found to reach the upper limit of the fatigue limit range of Ti–6Al–4V ELI by aging after severe cold rolling.
Maximum stress ;
994
105
106
107
108
Number of cycles to failure, Nf Fig. 5. S–N curves of Ti–6Al–4V ELI and SUS 316L in air and Ringer’s solution.
tent is reduced by degassing with nitrogen whereas it does not decrease without degassing. However, as shown in Fig. 8 [8], the fatigue strength of Ti–6Al–4V ELI obtained by uniaxial fatigue tests does not degrade in a living body environment, e.g., in the body of a living rabbit, which is in contrast to the fatigue strength of Ti–6Al–4V ELI in air. Therefore, there is a higher possibility of corrosion fatigue occurring under the bending fatigue condition in titanium alloys. In other words, the passive film formed on the surface of titanium alloys is considered to fracture more easily under the bending condition, and corrosion fatigue may
Fig. 4. S–N curves of TNTZST, TNTZCR, and TNTZST and TNTZCR conducted with aging at 598 K, 673 K and 723 K for 259.2 ks with those of Ti–6Al– 4V ELI and Ti–6Al–7Nb in air. TNTZST and TNTZCR indicate as-solutionized Ti–29Nb–13Ta–4.6Zr and as severe cold rolled Ti–29Nb–13Ta–4.6Zr, respectively.
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
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Stress, MPa
400
300
200
Csst Cr-Co-Mo alloy ( y = 420MPa) Air, R.T. Ringer’s solution, R.T.
100
106
105
10 4
107
Cycles to failure
400 : Air : Ringer’s solution : Ringer’s solution + N2 gas f=10 Hz
500 400 300 10 4
10 5 10 6 10 7 Number of cycles to failure, Nf
10 8
Maximum stress, σmax (MPa)
occur because it takes a relatively longer time for a fractured passive film to recover due to the low oxygen content. Fig. 9 [9] compares the result of the fatigue test of Co– Cr alloy in PBS with that in air. The fatigue strength of Co–Cr alloy in PBS is lower than in air in the low cycle fati-
1100
Ti-6Al-4V 10Hz
in air in rabbit
1000 900 800 not broken
700 600 3 10
10 4
10 5
10 6
Co-Cr alloy 300
200
Air f=20Hz PBS(-) f=2Hz 100
Fig. 7. S–N curves of Ti–5Al–2.5Fe conducted with annealing at 973 k for 7.2 ks followed by air cooling in air, Ringer’s solution, and Ringer’s solution with N2 gas.
1200
Stress amplitude, σa/MPa
(MPa)
600
Maximum stress ;
700
max
Fig. 6. Comparison of fatigue lives for cast material (type 1) in air and in simulated body solution.
10 7
Number of cycles to failure, Nf Fig. 8. S–N curves of Ti–6Al–4V obtained from uniaxial fatigue tests in air and rabbit.
10 4
10 5
10 6
10 7
10 8
10 9
Cycles to failure , Nf Fig. 9. S–N curves of Co–Cr alloy in the air and in PBS( ).
gue life region, while it is almost similar to that in air in the high cycle fatigue life region. 4. Bioactive surface modification and fatigue strength Among the metallic biomaterials, the biocompatibility of titanium alloys is the highest, but these alloys are not bioactive. Therefore, they are subjected to surface modification using bioactive ceramics in order to further improve their biocompatibility. There exist many processes for bioactive ceramic surface modification. They are roughly grouped into dry process, e.g., the spray method [10], and wet processes, e.g., the alkali method, and the dip-coating method. The fatigue characteristics of metallic biomaterials subjected to these different surface modification processes are also significant. For example, Fig. 10 [11] shows the results of the fatigue tests of Ti–29Nb–13Ta–4.6Zr coated with calcium phosphate invert-glass ceramic by the dip-coating method. In this method, a mixture of calcium phosphate invert-glass
996
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
Maximum cyclic stress (MPa)
1000 TNTZ TNTZ1 TNTZ2
800
f=10 Hz
600
400
200 10 3
10 4
10 5
10 6
10 7
blasting, which is a pretreatment applied to the surface of the substrate. Fig. 12 [13] shows the fatigue strength of SUS 316L stainless steel coated with alumina (Al2O3), which is not a bioactive ceramic, but is highly biocompatible, by the plasma spray method and evaluated in physiological saline. For comparison, it is also shows the fatigue strength of SUS 316L stainless steel without alumina coating. The fatigue strength of SUS 316L stainless steel with alumina coating is higher than that without the coating because the corrosion of the steel (with coating) is inhibited by the dense alumina coated layer.
Number of cycles to failure Fig. 10. S–N curves of as solutionized Ti–29Nb–13Ta–4.6Zr (TNTZ), glass-ceramic coated Ti–29Nb–13Ta–4.6Zr (TNTZ1), and glass-ceramic coated Ti–29Nb–13Ta–4.6Zr followed by aging (TNTZ2).
and distilled water is coated on the specimen by dipping the specimen into the mixture. The specimen is then fired in order to precipitate the calcium phosphate system ceramics. The firing temperature is above the b transus temperature of Ti–29Nb–13Ta–4.6Zr; thus the fatigue strength of Ti–29Nb–13Ta–4.6Zr cannot be kept high when the dipcoating method is employed. Therefore, it is necessary to finally age the specimen in order to improve the fatigue. As shown in Fig. 10, the fatigue strength of Ti–29Nb– 13Ta–4.6Zr coated with calcium phosphate invert-glass ceramic increases remarkably after aging. It is possible to inhibit cracking or exfoliation of the calcium phosphate invert glass-ceramic layer or cracking between the layer and the substrate by controlling the thickness of the layer [11,12]. Fig. 11 [11] shows an SEM fractograph of calcium phosphate invert glass-ceramic coated Ti–29Nb–13Ta– 4.6Zr after the fatigue test. Cracking or exfoliation of the layer cannot be observed on the SEM fractograph. It should be noted that the crack initiation site of this case is the pit formed on the surface of the substrate by sand
5. Fatigue crack propagation in air and simulated body environment In order to inhibit the catastrophic failure of biomaterials, it is necessary to understand the fatigue crack initiation and fatigue crack propagation characteristics. The crack propagation characteristics are important in arriving at fail safe design for in structural materials. It is considered that the catastrophic failure of these materials can be avoided by stopping the crack extension during the stable crack propagation stage, even if the crack has already been initiated. It is important to evaluate both long and short crack (small crack) propagation characteristics. The trends in propagation rates of long and short cracks are not always similar. Fig. 13 [14] shows the relationship between the fatigue crack propagation rate, da/dN, and the nominal cyclic stress intensity factor, DK, and that between da/dN and the effective cyclic stress intensity factor, DKeff, for Ti– 6Al–4V ELI with Widmansta¨tten a and equiaxed a structures, and SUS 316L stainless steel obtained in air. When da/dN is plotted against DK, it is observed to decrease in the following order: Ti–6Al–4V ELI with the Widmansta¨tten a structure P SUS 316L stainless steel P Ti–6Al–4V ELI with equiaxed a structure. On the other hand, when
Stress amplitude σa , MPa
280 260 240 220 200 f = 2 Hz in p.s.s
180 160 140 10 4
alumina-sprayed annealed
10 5
10 6
10 7
Number of cycles to failure, Nf Fig. 11. General view of fatigue fracture surface of glass-ceramic coated Ti–29Nb–13Ta–4.6Zr (TNTZ1) (N = 557862).
Fig. 12. S–N curves for annealed or alumina-sprayed SUS 316L stainless steels in physiological saline solution.
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
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Crack depth a,[mm]
0.04 0.05 0.1
0.4
0.06 0.1 0.07 0.1
0.5 1.0 0.5 0.7
SUS304 Small crack Room air, f =10Hz max =
10 -3
max =
200MPa 230MPa
da/dN is plotted against DKeff, the crack propagation rate of Ti–6Al–4V ELI with the Widmansta¨tten a structure is nearly the same as that of SUS 316L stainless steel. The crack propagation rate of Ti–6Al–4V ELI with the equiaxed a structure is the highest. However, the differences in the crack propagation rates among these materials become relatively smaller. In particular, the crack propagation rate of Ti–6Al–4V ELI with the Widmansta¨tten a structure approaches that of Ti–6Al–4V ELI with the equiaxed a structure. Therefore, the crack closure effect in Ti– 6Al–4V ELI with the Widmansta¨tten a structure is greater than that in Ti–6Al–4V ELI with the equiaxed a structure. The microstructure strongly affects the crack propagation rate in a + b type titanium alloys. In a representative a + b type titanium alloy, Ti–6Al–4V, the long crack propagation rate in the Widmansta¨tten a structure is, in general, lower than that in the equiaxed a structure. The main reason for this phenomenon is large crack deflection in the Widmansta¨tten a structure [5]. Therefore, the crack closure behavior in the Widmansta¨tten a structure is greater than that in the equiaxed a structure [5]. On the other hand, the short crack propagation rate in the equiaxed a structure is, in general, lower than that in the Widmansta¨tten a structure. The ratio of the short fatigue crack propagation life to the total fatigue life is fairly high. Therefore, improving resistance against short fatigue crack propagation is very effective in improving the total fatigue life [15,16]. It is reported that the fatigue crack propagation rates of SUS 304 stainless steel, which is also austenitic stainless steel, evaluated in air and 3% NaCl solution is clearly different in the very short fatigue crack region, and the short crack propagation rate in 3% NaCl solution is higher than that in air, as shown in Fig. 14 [17]. It is also reported that the long fatigue crack propagation rate of SUS 304 stainless steel in air is almost the same as that in 3% NaCl solution, as shown in Fig. 15 [17]; however, in the range of high stress intensity factor, the crack propagation rate in 3% NaCl solution is slightly higher than that in air.
max =
200MPa
10 -4
10 -5
10 -6 Long crack Room air 3%NaCl sol.
10 -7 1
5
K
10
K elf
20
30
Stress intensity factor range K MPa m Fig. 14. da/dN–DK relationships for small cracks in SUS 304.
10-3 Long crack
Crack growth rate da/dN mm/cycle
Fig. 13. The fatigue crack propagation rate, da/dN, as a function of the nominal cyclic stress intensity factor range, DK, and as a function of the effective cyclic stress intensity factor range, DKeff, in the case of variously heat-treated Ti–6Al–4V ELI and SUS 316 L.
Crack growth rate
da/dN
mm/cycle
3%NaCl sol., f =1Hz
ΔK ΔKeff SUS304 Room air 3%NaCl sol.
10-4
10-5
10-6
10-7
1
10 5 20 30 40 Stress intensity factor range ΔK
ΔKeff MPa m
Fig. 15. da/dN–DK and DKeff relationships for long cracks.
It has been reported that for Co–Cr alloy, the fatigue crack propagation rate in Ringer’s solution is higher than that in air [18]. When da/dN is plotted against DK, the long fatigue crack propagation rates of Ti–6Al–4V ELI and Ti–5Al– 2.5Fe in Ringer’s solution are higher than those in air as
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
shown in Fig. 16, but they are nearly the same in air and Ringer’s solution when da/dN is plotted against DKeff as shown in Fig. 17. In this case, it is reported that the crack closure reduces in Ringer’s solution because the number of secondary cracks and the fatigue fracture surface in Ringer’s solution are smaller than those in air. This suggests that the fatigue fracture surface corrodes and dissolves in Ringer’s solution.
10-6
10-7
da/dN , m/cycle
998
10-8
10-9
6. Fretting fatigue strength in air and simulated body environment
Ti-5Al-2.5Fe B(Equiaxed α + fine precipitated α) Ti-6Al-4V ELI F(Widmanstätten α)
10-10 1
There is a high possibility for fretting fatigue, in which friction wear and cyclic loading overlap, to occur between two bodies, such as between a plate and screw, or between a stem and bone. This may cause a failure of the implants. Therefore, it is essential to determine the fretting fatigue characteristics of metallic biomaterials. The fretting fatigue strength of these materials is known to decrease by 1/2–1/3 of the plain fatigue strength [19,20]. 10-6 Ti-5Al-2.5Fe B(Equiaxed α + fine precipitated α) Ti-6Al-4V ELI F(Widmanstätten α)
da/dN, m/cycle
10-7
in Air in Ringer’s solution
10-8
10-9
10-10 1
10
100
ΔK , MPa m1/2
Fig. 16. Fatigue crack propagation rate, da/dN, as a function of nominal cyclic stress intensity factor range, DK, in the case of Ti–5Al–2.5Fe with (equiaxed a + fine precipitated a) and Ti–6Al–4V ELI with Widmansta¨tten a in air and in Ringer’s solution.
10
in Air in Ringer’s solution
100
ΔKeff , MPa m1/2
Fig. 17. Fatigue crack propagation rate, da/dN, as a function of nominal cyclic stress intensity factor range, DKeff, in the case of Ti–5Al–2.5Fe with (equiaxed a + fine precipitated a) and Ti–6Al–4V ELI with Widmansta¨tten a in air and in Ringer’s solution.
Fig. 18 [19] compares the plain and the fretting fatigue limits of SUS 316L stainless steel, Co–Cr alloy, pure Ti (JIS Grade 2), annealed Ti–6Al–4V and aged Ti–6Al–4V in air and PBS ( ). For each material, the fretting fatigue limit is smaller than the plain fatigue limit in both air and PBS ( ). The ratio of the plain fatigue limit to the fretting fatigue limit is the lowest for Ti–6Al–4V, while it is the highest for Co–Cr alloy. This is considered to indicate the excellent friction wear resistance of Co–Cr alloy. The fretting fatigue limit of every material in PBS ( ), except Co–Cr alloy, is obviously smaller than that in air. This is because fretting causes the formation of pits and is considered to enhance the corrosion [20]. In the case of the plain fatigue and the fretting fatigue tests of Ti–29Nb–13Ta–4.6Zr in Ringer’s solution, the plain fatigue strength is lower than the fretting fatigue strength in the low cycle fatigue life region. On the other hand, the fretting fatigue strength is lower than the plain fatigue strength in the high cycle fatigue life region, as shown in Fig. 19 [20]. The lubrication influence of Ringer’s
Fig. 18. (a) Plain fatigue and (b) fretting fatigue limits of metallic biomaterials.
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
solution is greater than its corrosion influence in the low cycle fatigue life region, while, its corrosion influence is greater in the high cycle fatigue life region [20]. It has been reported that the decrease in fatigue strength due to fretting can be inhibited by decreasing Young’s modulus of the material, as shown in Fig. 20 [20]. The low modulus b type titanium alloy, which has recently gained considerable attention is advantageous from the viewpoint of inhibiting the decrease in fatigue strength due to fretting. 7. Fatigue strength of wire Wires composed of metallic biomaterials are useful in biomedical and dental applications, e.g., stents, guide wire of a catheter, surgical wire, and orthodontic wire. An interesting application of metallic wire is high nitrogen stainless steel wire used for the electrode of an FES (functional electrical stimulation). From the viewpoint of allergic reactions, low Ni stainless steel wire, that is, the high nitrogen stainless steel wire (22.0% Cr, 10% Ni,
999
6.02% Mn, 2% Mo, 0.41% N and bal. Fe) is preferred to SUS 316L stainless steel. In the case of wire, the rotating bending fatigue strength is important. Fig. 21 [21] shows the relationship between the number of rotations to failure, N, of electrode wires composed of SUS 316L stainless steel, NAS604PH (Co alloy), and NAS106N (high nitrogen stainless steel) and the distance between chucks, d, obtained from a dual-driven rotating–bending failure method in air. The rotating–bending fatigue strength of high nitrogen stainless steel is the highest. Titanium is expected to be a suitable material for this type of wire. Low modulus b type titanium alloys are also used for wires for stents, catheters, and orthopedic, surgical and orthodontic equipment. The shape memory alloy TiNi is widely used for catheters or orthodontic wires. However, TiNi contains a large amount of Ni, which has been reported to be an allergen and is also brittle. Therefore, Ni-free shape memory or super elastic titanium alloys with low moduli are being developed. As orthodontic wires, b type Ti–Mo–Zr–Sn has been put to practical use. Very recently, low modulus b type Ti–29Nb–13Ta–4.6Zr subjected to severe cold working and heat treatment has exhibited super elastic behavior as shown in Fig. 22 [22]. Fig. 23 [23] shows the notched fatigue strength of Ti–29Nb–13Ta– 4.6Zr wire with a diameter of 1.0 mm along with those of pure Ti, TiNi, and SUS 316L stainless steel wires with diameters of 1.0 mm. The notched fatigue strength of SUS 316L stainless steel wire is the highest in both the low and high cycle fatigue life regions. The notched fatigue strength of pure Ti wire is the lowest in the low cycle fatigue life region, and that of Ti–29Nb–13Ta–4.6Zr wire is slightly higher than that of TiNi wire. However, in the high cycle fatigue life region, the notched fatigue strengths of pure Ti, TiNi, and Ti–29Nb–13Ta–4.6Zr wires are nearly the same. In Japan, the official license for the application of Ti–29Nb–13Ta–4.6Zr for orthodontic wire has been issued very recently. Thus, Ti–29Nb–13Ta–4.6Zr wire will
Fig. 19. S–N curves of as-solutionized Ti–29Nb–13Ta–4.6Zr obtained from plain fatigue and fretting fatigue tests in air and Ringer’s solution.
: SUS316L
Ti-15Mo-5Zr-3Al (Annealing)
80
Ti-29Nb-13Ta-4.6Zr
70
STA
6
4
2
60 Ti-29Nb-13Ta-4.6Zr
50 1.0
No failure > 107
: NAS106N
N / 10 4
Modulus of elasticity, E / GPa
: NAS604PH 90
ST
0 1.5
2.0
2.5
3.0
3.5
4.0
Pf / Ff Fig. 20. Relationships between fretting damage ratio, Pf/Ff, modulus of elasticity; Pf: plain fatigue limit and Ff: fretting fatigue limit.
0
2
4
6
8
10
d / mm Fig. 21. Relationship between number of rotation of electrode materials to failure, N, and distance between chucks, d, by a dual-driven rotating– bending failure method in air.
1000
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000 700
Stress,MPa
600 500 400 300 200 100 0.5
0
1.0
2.0
1.5
2.5
3.0
3.5
4.0
Strain,% Fig. 22. Tensile loading–unloading stress–strain curves of drawn wire of Ti–29Nb–13Ta–4.6Zr with a diameter of 1.0 mm; total elastic strain: 2.7%.
TNTZd1.0 Pure Ti
800
Non Failure at 107 Cycles
Maximum Cyclic Stress, σmax / MPa
1000
Ti-Ni SUS316L
600
400
200 High Cycle Fatigue Life Region
Low Cycle Fatigue Life Region
0 10 3
10 4
10 5
10 6
10 7
10 8
Number of cycles to failure, Nf
Fig. 23. S–N curves of Ti–29Nb–13Ta–4.6Zr (TNTZd1) and pure Ti, Ti– Ni and SUS 316L stainless steel wires with a diameter of 1.0 mm obtained from notch-fatigue tests.
be put to practical use as orthodontic wire in the very near future. 8. Summarizing conclusions There exist a limited number of reports on the systematic investigation or comparison of fatigue characteristics, including fretting fatigue characteristics, with respect to related microstructures. Moreover, the number of reports on fatigue characteristics obtained in vitro or in vivo is very small. Further, the scatter of the fatigue data among the various studies is very large. Thus, the fatigue characteristics of metallic biomaterials should be researched further, and reliable fatigue data should be accumulated. References [1] Niinomi M. Mechanical properties of biomedical titanium alloys. Mater Sci Eng A 1998;243:231–6. [2] Kim JH, Niinomi M, Akahori T. Influence of bone structure on mechanical properties of bovine and swine compact bones. Mater Sci Forum 2005;475–479:2407–10.
[3] Breme HJ, Helsen JA. In: Helsen JA, Breme HJ, editors. Selection of materials. New York, USA: John Wiley and Sons; 1998. p. 1–36. [4] Akahori T, Niinomi M, Ishimizu K, Fukui H, Suzuki A. Effect of thermomehcanical treatment on fatigue characteristics of Ti–29NB– 13Ta–46Zr for biomedical applications. J Jpn Inst Met 2003;67:652–60. [5] Niinomi M, Kobayashi T, Akmoulin IA. Mechanical properties of titanium alloys. J Jpn Inst Light Met 1992;42:605–13. [6] Niinomi M, Kobayashi T, Toriyama O, Kawakami N, Ishida Y, Matsuyama Y. Fracture characteristics, microstructure, and tissue reaction of Ti–5Al–25Fe for orthopedic surgery. Metall Mater Trans A 1996;27A:3925–35. [7] Kumar P, Hickl AJ, Asphahani AI, Lawley A. Properties and characteristics of cast wrought and powder metallurgy (P/M) processed cobalt–chromium–molybdenum implant materials. ASTM STP 1985;859:30–56. [8] Niinomi M. Fatigue characteristics and microstructure of titanium alloys for biomedical applications. In: Proc Fatigue’2002, Blom AF, editors, EMAS 2002; 3/5. pp. 2073–83. [9] Maruyama N, Kobayashi T, Nakazawa K, Sumita M, Sato M. Fatigue and fretting behavior of Ni free Co–Cr alloy in a pseudobody fluid. J Jpn Biomater Soc 1999;17:172–9. [10] Niinomi M. Recent research and development in titanium alloys for biomedical applications and healthcare goods. Sci Technol Adv Mater 2003;4(4):445–54. [11] Li SJ, Niinomi M, Akahori T, Kasuga T, Yang R, Hao YL. Fatigue characteristics of bioactive glass-ceramic-coated Ti–29Nb–13Ta–46Zr for biomedical applications. Biomaterials 2004;25:3341–9. [12] Niinomi M, Akahori T, Yamaguchi T, Kasuga T, Fukui H, Suzuki A. Aging characteristics and mechanical properties of calcium phosphate invert glass-ceramic coated Ti–29Nb–13Ta–46Zr for biomedical applications. J Jpn Inst Met 2003;67:604–13. [13] Kisuki H, Sugano M. Fatigue properties of Al2O3 plasma-sprayed SUS316L stainless steel in physiological saline solution. J High Temp Suppl 1998;24:227–34. [14] Niinomi M, Saga A, Fukunaga K. Long crack growth behavior of implant materials Ti–5Al–25Fe in air and simulated body environment with relating microstructure. Int J Fatigue 2000;22: 887–97. [15] Niinomi M. Fatigue. J Iron Steel Inst Jpn 2000;5:82–8. [16] Akahori T, Niinomi M, Fukunaga K, Inagaki I. Effects of microstructure on the short fatigue crack initiation and propagation characteristics of biomedical a/b titanium alloys. Metall Mater Trans A 2000;31A:1949–58. [17] Nakajima M, Tokaji K, Akatsuka Y. Corrosion fatigue crack initiation and early growth behavior in stainless steel. J Jpn Soc Eng A 1997;63:487–92. [18] Hanawa T, Nakazaw K, Maruyama N, Yamamoto A, Hiromoto S, Sumita M, et al. Study on structural biomaterials with excellent biocompatibility. Res Rep Natl Inst Met 2000;22:223–39. [19] Hanawa T. Development of metallic biomaterials, Mater Jpn, Present state and future strategy 2003; 43. pp. 176–81. [20] Niinomi M, Akahori T, Yabunaka T, Fukui H, Suzuki A. Fretting fatigue characteristics of b type titanium alloy for biomedical applications in air and simulated body environment. J Iron Steel Inst Jpn 2003;88:553–60. [21] Iguchi Y. Development of the implanted stainless steel electrode for functional electrical stimulation based on ultra clean refining processes. Bull Iron Steel Inst Jpn 1999;4:737–40. [22] Niinomi M. Recent research and development in titanium alloys for biomedical applications and healthcare goods. STAM 2003;4: 445–54. [23] Akahori T, Niinomi M, Toda H, Yamauchi K, Fukui H, Ogawa M. Mechanical properties and deformation behavior of b-type titanium alloy (TNTZ) drawn wires. J Jpn Inst Met 2005;69(7):530–7.
International Journal of Fatigue 29 (2007) 992–1000
www.elsevier.com/locate/ijfatigue
Fatigue characteristics of metallic biomaterials M. Niinomi
*
Department of Biomaterials Science, Institute for Materials Research, Tohoku University, 2-1-1, Katahira Aoba-ku, Sendai, Miyagi 980-8577, Japan Received 16 November 2005; received in revised form 30 August 2006; accepted 26 September 2006 Available online 13 November 2006
Abstract It is essential to determine the fatigue characteristics of metallic biomaterials for long-term safe usage of implants in living body such as hip joints, bone plates, and dental implants. On the basis of reported data, the fatigue strength, fretting fatigue strength, and fatigue crack propagation rate of metallic biomaterials – mainly titanium alloys, Co alloys, and stainless steels – in air, a living body and a simulated body environment are described extensively with respect to related microstructures. Further, the fatigue strength of titanium alloys coated with bioactive ceramic and that of biometallic wire are examined. Ó 2006 Elsevier Ltd. All rights reserved. Keywords: Fatigue; Fatigue crack; Fretting fatigue; Metallic biomaterials; Titanium alloy; Co alloy; Stainless steel
1. Introduction Implants that function as bones – for example, implants, which replace failed hard tissue, such as artificial hip joints, artificial joints, bone plates, and dental implants – are usually used under severe cyclic loading conditions. Therefore, metallic materials that typically exhibit high strength, ductility and toughness are the main candidates for the structural biomaterials of these implants. Further, the abovementioned implants must exhibit high biocompatibility, high performance, and reliability for long-term use. The reliability of the implants after implantations is determined by their fracture and wear after the critical period of infection. Therefore, the mechanical properties of structural biomaterials in a living body environment – such as fatigue, toughness, and wear resistance – need to be evaluated and then improved in order to confidently use the implants for a long period of time. With regard to the fracture of structural biomaterials, fatigue fracture occurs occasionally; it is considered to be a crucial problem among the various types of fractures.
*
Tel.: +81 22 215 2574; fax: +81 22 215 2533. E-mail address: [email protected]
0142-1123/$ - see front matter Ó 2006 Elsevier Ltd. All rights reserved. doi:10.1016/j.ijfatigue.2006.09.021
Fatigue with fretting, i.e., fretting fatigue, is a type of fatigue that can occur between two bodies, such as between a bone plate and screw. Fatigue characteristics are closely related with the microstructures. The microstructures in metallic structural biomaterials change according to the processing and heat treatment employed. Therefore, in this paper, the fatigue characteristics of metallic structural biomaterials are described extensively with respect to related microstructures or processing for microstructural control. 2. Fatigue strength in air Fig. 1 [1] shows an example of the fatigue limits of stainless steels, Co alloy, and titanium and its alloys as representative metallic biomaterials in air; the fatigue limits are scattered depending on factors such the fabricating process, surface condition, microstructure, and fatigue condition. The fatigue limit of bovine bone [2] is also shown in the Fig. 1. This limit decreases in the following order: Co alloy P Ti–6Al–4V P 316L stainless steel. However, the limit of each metallic biomaterial shows a fairly large scatter due to the abovementioned factors. The fatigue limit of each metallic biomaterial is higher than that of bovine bone. Figs. 2 and 3 [3,4] show the relationship between the rotating bending fatigue limit and the elongation up to fracture and that between the fatigue strength and Young’s
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
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Fig. 1. Fatigue strength at 107 cycles of biomedical stainless steel, Co alloys, titanium and it’s alloys, and bone. Data without designation of rotating bending are those obtained from uniaxial fatigue tests.
Fig. 2. Elongation at fracture as a function of the fatigue strength of metallic biomaterials.
modulus for various metallic biomaterials, respectively. The rotating bending fatigue strength increases with the elongation up to fracture. The rotating bending fatigue strength is relatively higher and Young’s modulus is relatively low in a + b titanium alloys. Therefore, the BF value of titanium alloys, which is obtained by dividing the fatigue strength with Young’s modulus, is high (.5.2). The BF values of materials such as 316L stainless steel, Co–Cr alloy, Co–Ni–Cr alloy, pure Ti, and pure Nb, lie between 1.2 and 2.3 [3]. This indicates that the mechanical biocompatibility of titanium alloys is excellent in comparison with that of other metallic biomaterials. The BF value of aged Ti–29Nb–13Ta–4.6Zr (TNTZ alloy), which has recently been proposed for use in biomedical applications, is very high (.9.6) [4].
Fig. 3. Fatigue strength and Young’s modulus of each metallic biomaterial. TNTZ means low modulus b type titanium alloy, Ti–29Nb–13Ta– 4.6Zr.
Among the titanium alloys that have received considerable attention as metallic biomaterials, the fatigue limit of pure Ti, which is a type alloy, is the lowest. The fatigue limits of Ti–6Al–4V and Ti–6Al–7Nb, which are a + b type alloys, are among the highest. The fatigue limit of solutionized b type alloy is low; however, it increases drastically to approximate that of an a + b type alloy by aging after solutionizing due to precipitation of the a phase. The fatigue strength of a + b and b type alloys varies considerably with
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
Air Ringer’s, f=10 Hz
Ti-6Al-4V ELI : A (Equiaxed ) : C (Windmanstatten
700
(MPa)
600 500
400
300
200
100 104
3. Fatigue strength in simulated body environment In order to estimate the fatigue strength of metallic biomaterials in a living body environment, it is essential to evaluate it in a simulated body environment. Fig. 5 [6] compares the rotating bending fatigue strength of Ti–6Al–4V ELI and SUS 316L stainless steel in Ringer’s solution with those in air. The fatigue strength of Ti–6Al– 4V ELI in Ringer’s solution is similar to that in air, while that of SUS 316L stainless steel in Ringer’s solution is lower than that in air. Therefore, corrosion fatigue occurs in Ringer’s solution for SUS 316L stainless steel. The fatigue strength of Co–Cr–Mo is also reported to decrease in Ringer’s solution as shown in Fig. 6 [7]. As shown in Fig. 7 [6], the fatigue strength of Ti–5Al– 2.5Fe in Ringer’s solution decreases when the oxygen con-
)
: SUS 316L
max
the microstructure. In a + b type alloys, the fatigue strength of an equiaxed a structure is, in general, higher than that of an acicular or Widmansta¨tten a structure [5]. It is possible to improve the fatigue strength of b type alloys by thermomechanical treatment involving of cold working and heat treatment. Cold working can be performed very easily in b type alloys. Fig. 4 [4] shows the fatigue strength of low modulus Ti–29Nb–13Ta–4.6Zr – which was obtained after solutionizing, aging (TNTZST aged at 598 K and TNTZST aged at 673 K), severe cold rolling and aging (TNTZCR aged at 598 K and TNTZCR aged at 673 K) – with the fatigue limits of Ti–6Al–4V ELI and Ti–6Al–7Nb. The fatigue limit of Ti–29Nb–13Ta–4.6Zr is found to reach the upper limit of the fatigue limit range of Ti–6Al–4V ELI by aging after severe cold rolling.
Maximum stress ;
994
105
106
107
108
Number of cycles to failure, Nf Fig. 5. S–N curves of Ti–6Al–4V ELI and SUS 316L in air and Ringer’s solution.
tent is reduced by degassing with nitrogen whereas it does not decrease without degassing. However, as shown in Fig. 8 [8], the fatigue strength of Ti–6Al–4V ELI obtained by uniaxial fatigue tests does not degrade in a living body environment, e.g., in the body of a living rabbit, which is in contrast to the fatigue strength of Ti–6Al–4V ELI in air. Therefore, there is a higher possibility of corrosion fatigue occurring under the bending fatigue condition in titanium alloys. In other words, the passive film formed on the surface of titanium alloys is considered to fracture more easily under the bending condition, and corrosion fatigue may
Fig. 4. S–N curves of TNTZST, TNTZCR, and TNTZST and TNTZCR conducted with aging at 598 K, 673 K and 723 K for 259.2 ks with those of Ti–6Al– 4V ELI and Ti–6Al–7Nb in air. TNTZST and TNTZCR indicate as-solutionized Ti–29Nb–13Ta–4.6Zr and as severe cold rolled Ti–29Nb–13Ta–4.6Zr, respectively.
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
995
Stress, MPa
400
300
200
Csst Cr-Co-Mo alloy ( y = 420MPa) Air, R.T. Ringer’s solution, R.T.
100
106
105
10 4
107
Cycles to failure
400 : Air : Ringer’s solution : Ringer’s solution + N2 gas f=10 Hz
500 400 300 10 4
10 5 10 6 10 7 Number of cycles to failure, Nf
10 8
Maximum stress, σmax (MPa)
occur because it takes a relatively longer time for a fractured passive film to recover due to the low oxygen content. Fig. 9 [9] compares the result of the fatigue test of Co– Cr alloy in PBS with that in air. The fatigue strength of Co–Cr alloy in PBS is lower than in air in the low cycle fati-
1100
Ti-6Al-4V 10Hz
in air in rabbit
1000 900 800 not broken
700 600 3 10
10 4
10 5
10 6
Co-Cr alloy 300
200
Air f=20Hz PBS(-) f=2Hz 100
Fig. 7. S–N curves of Ti–5Al–2.5Fe conducted with annealing at 973 k for 7.2 ks followed by air cooling in air, Ringer’s solution, and Ringer’s solution with N2 gas.
1200
Stress amplitude, σa/MPa
(MPa)
600
Maximum stress ;
700
max
Fig. 6. Comparison of fatigue lives for cast material (type 1) in air and in simulated body solution.
10 7
Number of cycles to failure, Nf Fig. 8. S–N curves of Ti–6Al–4V obtained from uniaxial fatigue tests in air and rabbit.
10 4
10 5
10 6
10 7
10 8
10 9
Cycles to failure , Nf Fig. 9. S–N curves of Co–Cr alloy in the air and in PBS( ).
gue life region, while it is almost similar to that in air in the high cycle fatigue life region. 4. Bioactive surface modification and fatigue strength Among the metallic biomaterials, the biocompatibility of titanium alloys is the highest, but these alloys are not bioactive. Therefore, they are subjected to surface modification using bioactive ceramics in order to further improve their biocompatibility. There exist many processes for bioactive ceramic surface modification. They are roughly grouped into dry process, e.g., the spray method [10], and wet processes, e.g., the alkali method, and the dip-coating method. The fatigue characteristics of metallic biomaterials subjected to these different surface modification processes are also significant. For example, Fig. 10 [11] shows the results of the fatigue tests of Ti–29Nb–13Ta–4.6Zr coated with calcium phosphate invert-glass ceramic by the dip-coating method. In this method, a mixture of calcium phosphate invert-glass
996
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
Maximum cyclic stress (MPa)
1000 TNTZ TNTZ1 TNTZ2
800
f=10 Hz
600
400
200 10 3
10 4
10 5
10 6
10 7
blasting, which is a pretreatment applied to the surface of the substrate. Fig. 12 [13] shows the fatigue strength of SUS 316L stainless steel coated with alumina (Al2O3), which is not a bioactive ceramic, but is highly biocompatible, by the plasma spray method and evaluated in physiological saline. For comparison, it is also shows the fatigue strength of SUS 316L stainless steel without alumina coating. The fatigue strength of SUS 316L stainless steel with alumina coating is higher than that without the coating because the corrosion of the steel (with coating) is inhibited by the dense alumina coated layer.
Number of cycles to failure Fig. 10. S–N curves of as solutionized Ti–29Nb–13Ta–4.6Zr (TNTZ), glass-ceramic coated Ti–29Nb–13Ta–4.6Zr (TNTZ1), and glass-ceramic coated Ti–29Nb–13Ta–4.6Zr followed by aging (TNTZ2).
and distilled water is coated on the specimen by dipping the specimen into the mixture. The specimen is then fired in order to precipitate the calcium phosphate system ceramics. The firing temperature is above the b transus temperature of Ti–29Nb–13Ta–4.6Zr; thus the fatigue strength of Ti–29Nb–13Ta–4.6Zr cannot be kept high when the dipcoating method is employed. Therefore, it is necessary to finally age the specimen in order to improve the fatigue. As shown in Fig. 10, the fatigue strength of Ti–29Nb– 13Ta–4.6Zr coated with calcium phosphate invert-glass ceramic increases remarkably after aging. It is possible to inhibit cracking or exfoliation of the calcium phosphate invert glass-ceramic layer or cracking between the layer and the substrate by controlling the thickness of the layer [11,12]. Fig. 11 [11] shows an SEM fractograph of calcium phosphate invert glass-ceramic coated Ti–29Nb–13Ta– 4.6Zr after the fatigue test. Cracking or exfoliation of the layer cannot be observed on the SEM fractograph. It should be noted that the crack initiation site of this case is the pit formed on the surface of the substrate by sand
5. Fatigue crack propagation in air and simulated body environment In order to inhibit the catastrophic failure of biomaterials, it is necessary to understand the fatigue crack initiation and fatigue crack propagation characteristics. The crack propagation characteristics are important in arriving at fail safe design for in structural materials. It is considered that the catastrophic failure of these materials can be avoided by stopping the crack extension during the stable crack propagation stage, even if the crack has already been initiated. It is important to evaluate both long and short crack (small crack) propagation characteristics. The trends in propagation rates of long and short cracks are not always similar. Fig. 13 [14] shows the relationship between the fatigue crack propagation rate, da/dN, and the nominal cyclic stress intensity factor, DK, and that between da/dN and the effective cyclic stress intensity factor, DKeff, for Ti– 6Al–4V ELI with Widmansta¨tten a and equiaxed a structures, and SUS 316L stainless steel obtained in air. When da/dN is plotted against DK, it is observed to decrease in the following order: Ti–6Al–4V ELI with the Widmansta¨tten a structure P SUS 316L stainless steel P Ti–6Al–4V ELI with equiaxed a structure. On the other hand, when
Stress amplitude σa , MPa
280 260 240 220 200 f = 2 Hz in p.s.s
180 160 140 10 4
alumina-sprayed annealed
10 5
10 6
10 7
Number of cycles to failure, Nf Fig. 11. General view of fatigue fracture surface of glass-ceramic coated Ti–29Nb–13Ta–4.6Zr (TNTZ1) (N = 557862).
Fig. 12. S–N curves for annealed or alumina-sprayed SUS 316L stainless steels in physiological saline solution.
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
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Crack depth a,[mm]
0.04 0.05 0.1
0.4
0.06 0.1 0.07 0.1
0.5 1.0 0.5 0.7
SUS304 Small crack Room air, f =10Hz max =
10 -3
max =
200MPa 230MPa
da/dN is plotted against DKeff, the crack propagation rate of Ti–6Al–4V ELI with the Widmansta¨tten a structure is nearly the same as that of SUS 316L stainless steel. The crack propagation rate of Ti–6Al–4V ELI with the equiaxed a structure is the highest. However, the differences in the crack propagation rates among these materials become relatively smaller. In particular, the crack propagation rate of Ti–6Al–4V ELI with the Widmansta¨tten a structure approaches that of Ti–6Al–4V ELI with the equiaxed a structure. Therefore, the crack closure effect in Ti– 6Al–4V ELI with the Widmansta¨tten a structure is greater than that in Ti–6Al–4V ELI with the equiaxed a structure. The microstructure strongly affects the crack propagation rate in a + b type titanium alloys. In a representative a + b type titanium alloy, Ti–6Al–4V, the long crack propagation rate in the Widmansta¨tten a structure is, in general, lower than that in the equiaxed a structure. The main reason for this phenomenon is large crack deflection in the Widmansta¨tten a structure [5]. Therefore, the crack closure behavior in the Widmansta¨tten a structure is greater than that in the equiaxed a structure [5]. On the other hand, the short crack propagation rate in the equiaxed a structure is, in general, lower than that in the Widmansta¨tten a structure. The ratio of the short fatigue crack propagation life to the total fatigue life is fairly high. Therefore, improving resistance against short fatigue crack propagation is very effective in improving the total fatigue life [15,16]. It is reported that the fatigue crack propagation rates of SUS 304 stainless steel, which is also austenitic stainless steel, evaluated in air and 3% NaCl solution is clearly different in the very short fatigue crack region, and the short crack propagation rate in 3% NaCl solution is higher than that in air, as shown in Fig. 14 [17]. It is also reported that the long fatigue crack propagation rate of SUS 304 stainless steel in air is almost the same as that in 3% NaCl solution, as shown in Fig. 15 [17]; however, in the range of high stress intensity factor, the crack propagation rate in 3% NaCl solution is slightly higher than that in air.
max =
200MPa
10 -4
10 -5
10 -6 Long crack Room air 3%NaCl sol.
10 -7 1
5
K
10
K elf
20
30
Stress intensity factor range K MPa m Fig. 14. da/dN–DK relationships for small cracks in SUS 304.
10-3 Long crack
Crack growth rate da/dN mm/cycle
Fig. 13. The fatigue crack propagation rate, da/dN, as a function of the nominal cyclic stress intensity factor range, DK, and as a function of the effective cyclic stress intensity factor range, DKeff, in the case of variously heat-treated Ti–6Al–4V ELI and SUS 316 L.
Crack growth rate
da/dN
mm/cycle
3%NaCl sol., f =1Hz
ΔK ΔKeff SUS304 Room air 3%NaCl sol.
10-4
10-5
10-6
10-7
1
10 5 20 30 40 Stress intensity factor range ΔK
ΔKeff MPa m
Fig. 15. da/dN–DK and DKeff relationships for long cracks.
It has been reported that for Co–Cr alloy, the fatigue crack propagation rate in Ringer’s solution is higher than that in air [18]. When da/dN is plotted against DK, the long fatigue crack propagation rates of Ti–6Al–4V ELI and Ti–5Al– 2.5Fe in Ringer’s solution are higher than those in air as
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
shown in Fig. 16, but they are nearly the same in air and Ringer’s solution when da/dN is plotted against DKeff as shown in Fig. 17. In this case, it is reported that the crack closure reduces in Ringer’s solution because the number of secondary cracks and the fatigue fracture surface in Ringer’s solution are smaller than those in air. This suggests that the fatigue fracture surface corrodes and dissolves in Ringer’s solution.
10-6
10-7
da/dN , m/cycle
998
10-8
10-9
6. Fretting fatigue strength in air and simulated body environment
Ti-5Al-2.5Fe B(Equiaxed α + fine precipitated α) Ti-6Al-4V ELI F(Widmanstätten α)
10-10 1
There is a high possibility for fretting fatigue, in which friction wear and cyclic loading overlap, to occur between two bodies, such as between a plate and screw, or between a stem and bone. This may cause a failure of the implants. Therefore, it is essential to determine the fretting fatigue characteristics of metallic biomaterials. The fretting fatigue strength of these materials is known to decrease by 1/2–1/3 of the plain fatigue strength [19,20]. 10-6 Ti-5Al-2.5Fe B(Equiaxed α + fine precipitated α) Ti-6Al-4V ELI F(Widmanstätten α)
da/dN, m/cycle
10-7
in Air in Ringer’s solution
10-8
10-9
10-10 1
10
100
ΔK , MPa m1/2
Fig. 16. Fatigue crack propagation rate, da/dN, as a function of nominal cyclic stress intensity factor range, DK, in the case of Ti–5Al–2.5Fe with (equiaxed a + fine precipitated a) and Ti–6Al–4V ELI with Widmansta¨tten a in air and in Ringer’s solution.
10
in Air in Ringer’s solution
100
ΔKeff , MPa m1/2
Fig. 17. Fatigue crack propagation rate, da/dN, as a function of nominal cyclic stress intensity factor range, DKeff, in the case of Ti–5Al–2.5Fe with (equiaxed a + fine precipitated a) and Ti–6Al–4V ELI with Widmansta¨tten a in air and in Ringer’s solution.
Fig. 18 [19] compares the plain and the fretting fatigue limits of SUS 316L stainless steel, Co–Cr alloy, pure Ti (JIS Grade 2), annealed Ti–6Al–4V and aged Ti–6Al–4V in air and PBS ( ). For each material, the fretting fatigue limit is smaller than the plain fatigue limit in both air and PBS ( ). The ratio of the plain fatigue limit to the fretting fatigue limit is the lowest for Ti–6Al–4V, while it is the highest for Co–Cr alloy. This is considered to indicate the excellent friction wear resistance of Co–Cr alloy. The fretting fatigue limit of every material in PBS ( ), except Co–Cr alloy, is obviously smaller than that in air. This is because fretting causes the formation of pits and is considered to enhance the corrosion [20]. In the case of the plain fatigue and the fretting fatigue tests of Ti–29Nb–13Ta–4.6Zr in Ringer’s solution, the plain fatigue strength is lower than the fretting fatigue strength in the low cycle fatigue life region. On the other hand, the fretting fatigue strength is lower than the plain fatigue strength in the high cycle fatigue life region, as shown in Fig. 19 [20]. The lubrication influence of Ringer’s
Fig. 18. (a) Plain fatigue and (b) fretting fatigue limits of metallic biomaterials.
M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000
solution is greater than its corrosion influence in the low cycle fatigue life region, while, its corrosion influence is greater in the high cycle fatigue life region [20]. It has been reported that the decrease in fatigue strength due to fretting can be inhibited by decreasing Young’s modulus of the material, as shown in Fig. 20 [20]. The low modulus b type titanium alloy, which has recently gained considerable attention is advantageous from the viewpoint of inhibiting the decrease in fatigue strength due to fretting. 7. Fatigue strength of wire Wires composed of metallic biomaterials are useful in biomedical and dental applications, e.g., stents, guide wire of a catheter, surgical wire, and orthodontic wire. An interesting application of metallic wire is high nitrogen stainless steel wire used for the electrode of an FES (functional electrical stimulation). From the viewpoint of allergic reactions, low Ni stainless steel wire, that is, the high nitrogen stainless steel wire (22.0% Cr, 10% Ni,
999
6.02% Mn, 2% Mo, 0.41% N and bal. Fe) is preferred to SUS 316L stainless steel. In the case of wire, the rotating bending fatigue strength is important. Fig. 21 [21] shows the relationship between the number of rotations to failure, N, of electrode wires composed of SUS 316L stainless steel, NAS604PH (Co alloy), and NAS106N (high nitrogen stainless steel) and the distance between chucks, d, obtained from a dual-driven rotating–bending failure method in air. The rotating–bending fatigue strength of high nitrogen stainless steel is the highest. Titanium is expected to be a suitable material for this type of wire. Low modulus b type titanium alloys are also used for wires for stents, catheters, and orthopedic, surgical and orthodontic equipment. The shape memory alloy TiNi is widely used for catheters or orthodontic wires. However, TiNi contains a large amount of Ni, which has been reported to be an allergen and is also brittle. Therefore, Ni-free shape memory or super elastic titanium alloys with low moduli are being developed. As orthodontic wires, b type Ti–Mo–Zr–Sn has been put to practical use. Very recently, low modulus b type Ti–29Nb–13Ta–4.6Zr subjected to severe cold working and heat treatment has exhibited super elastic behavior as shown in Fig. 22 [22]. Fig. 23 [23] shows the notched fatigue strength of Ti–29Nb–13Ta– 4.6Zr wire with a diameter of 1.0 mm along with those of pure Ti, TiNi, and SUS 316L stainless steel wires with diameters of 1.0 mm. The notched fatigue strength of SUS 316L stainless steel wire is the highest in both the low and high cycle fatigue life regions. The notched fatigue strength of pure Ti wire is the lowest in the low cycle fatigue life region, and that of Ti–29Nb–13Ta–4.6Zr wire is slightly higher than that of TiNi wire. However, in the high cycle fatigue life region, the notched fatigue strengths of pure Ti, TiNi, and Ti–29Nb–13Ta–4.6Zr wires are nearly the same. In Japan, the official license for the application of Ti–29Nb–13Ta–4.6Zr for orthodontic wire has been issued very recently. Thus, Ti–29Nb–13Ta–4.6Zr wire will
Fig. 19. S–N curves of as-solutionized Ti–29Nb–13Ta–4.6Zr obtained from plain fatigue and fretting fatigue tests in air and Ringer’s solution.
: SUS316L
Ti-15Mo-5Zr-3Al (Annealing)
80
Ti-29Nb-13Ta-4.6Zr
70
STA
6
4
2
60 Ti-29Nb-13Ta-4.6Zr
50 1.0
No failure > 107
: NAS106N
N / 10 4
Modulus of elasticity, E / GPa
: NAS604PH 90
ST
0 1.5
2.0
2.5
3.0
3.5
4.0
Pf / Ff Fig. 20. Relationships between fretting damage ratio, Pf/Ff, modulus of elasticity; Pf: plain fatigue limit and Ff: fretting fatigue limit.
0
2
4
6
8
10
d / mm Fig. 21. Relationship between number of rotation of electrode materials to failure, N, and distance between chucks, d, by a dual-driven rotating– bending failure method in air.
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M. Niinomi / International Journal of Fatigue 29 (2007) 992–1000 700
Stress,MPa
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0
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Fig. 23. S–N curves of Ti–29Nb–13Ta–4.6Zr (TNTZd1) and pure Ti, Ti– Ni and SUS 316L stainless steel wires with a diameter of 1.0 mm obtained from notch-fatigue tests.
be put to practical use as orthodontic wire in the very near future. 8. Summarizing conclusions There exist a limited number of reports on the systematic investigation or comparison of fatigue characteristics, including fretting fatigue characteristics, with respect to related microstructures. Moreover, the number of reports on fatigue characteristics obtained in vitro or in vivo is very small. Further, the scatter of the fatigue data among the various studies is very large. Thus, the fatigue characteristics of metallic biomaterials should be researched further, and reliable fatigue data should be accumulated. References [1] Niinomi M. Mechanical properties of biomedical titanium alloys. Mater Sci Eng A 1998;243:231–6. [2] Kim JH, Niinomi M, Akahori T. Influence of bone structure on mechanical properties of bovine and swine compact bones. Mater Sci Forum 2005;475–479:2407–10.
[3] Breme HJ, Helsen JA. In: Helsen JA, Breme HJ, editors. Selection of materials. New York, USA: John Wiley and Sons; 1998. p. 1–36. [4] Akahori T, Niinomi M, Ishimizu K, Fukui H, Suzuki A. Effect of thermomehcanical treatment on fatigue characteristics of Ti–29NB– 13Ta–46Zr for biomedical applications. J Jpn Inst Met 2003;67:652–60. [5] Niinomi M, Kobayashi T, Akmoulin IA. Mechanical properties of titanium alloys. J Jpn Inst Light Met 1992;42:605–13. [6] Niinomi M, Kobayashi T, Toriyama O, Kawakami N, Ishida Y, Matsuyama Y. Fracture characteristics, microstructure, and tissue reaction of Ti–5Al–25Fe for orthopedic surgery. Metall Mater Trans A 1996;27A:3925–35. [7] Kumar P, Hickl AJ, Asphahani AI, Lawley A. Properties and characteristics of cast wrought and powder metallurgy (P/M) processed cobalt–chromium–molybdenum implant materials. ASTM STP 1985;859:30–56. [8] Niinomi M. Fatigue characteristics and microstructure of titanium alloys for biomedical applications. In: Proc Fatigue’2002, Blom AF, editors, EMAS 2002; 3/5. pp. 2073–83. [9] Maruyama N, Kobayashi T, Nakazawa K, Sumita M, Sato M. Fatigue and fretting behavior of Ni free Co–Cr alloy in a pseudobody fluid. J Jpn Biomater Soc 1999;17:172–9. [10] Niinomi M. Recent research and development in titanium alloys for biomedical applications and healthcare goods. Sci Technol Adv Mater 2003;4(4):445–54. [11] Li SJ, Niinomi M, Akahori T, Kasuga T, Yang R, Hao YL. Fatigue characteristics of bioactive glass-ceramic-coated Ti–29Nb–13Ta–46Zr for biomedical applications. Biomaterials 2004;25:3341–9. [12] Niinomi M, Akahori T, Yamaguchi T, Kasuga T, Fukui H, Suzuki A. Aging characteristics and mechanical properties of calcium phosphate invert glass-ceramic coated Ti–29Nb–13Ta–46Zr for biomedical applications. J Jpn Inst Met 2003;67:604–13. [13] Kisuki H, Sugano M. Fatigue properties of Al2O3 plasma-sprayed SUS316L stainless steel in physiological saline solution. J High Temp Suppl 1998;24:227–34. [14] Niinomi M, Saga A, Fukunaga K. Long crack growth behavior of implant materials Ti–5Al–25Fe in air and simulated body environment with relating microstructure. Int J Fatigue 2000;22: 887–97. [15] Niinomi M. Fatigue. J Iron Steel Inst Jpn 2000;5:82–8. [16] Akahori T, Niinomi M, Fukunaga K, Inagaki I. Effects of microstructure on the short fatigue crack initiation and propagation characteristics of biomedical a/b titanium alloys. Metall Mater Trans A 2000;31A:1949–58. [17] Nakajima M, Tokaji K, Akatsuka Y. Corrosion fatigue crack initiation and early growth behavior in stainless steel. J Jpn Soc Eng A 1997;63:487–92. [18] Hanawa T, Nakazaw K, Maruyama N, Yamamoto A, Hiromoto S, Sumita M, et al. Study on structural biomaterials with excellent biocompatibility. Res Rep Natl Inst Met 2000;22:223–39. [19] Hanawa T. Development of metallic biomaterials, Mater Jpn, Present state and future strategy 2003; 43. pp. 176–81. [20] Niinomi M, Akahori T, Yabunaka T, Fukui H, Suzuki A. Fretting fatigue characteristics of b type titanium alloy for biomedical applications in air and simulated body environment. J Iron Steel Inst Jpn 2003;88:553–60. [21] Iguchi Y. Development of the implanted stainless steel electrode for functional electrical stimulation based on ultra clean refining processes. Bull Iron Steel Inst Jpn 1999;4:737–40. [22] Niinomi M. Recent research and development in titanium alloys for biomedical applications and healthcare goods. STAM 2003;4: 445–54. [23] Akahori T, Niinomi M, Toda H, Yamauchi K, Fukui H, Ogawa M. Mechanical properties and deformation behavior of b-type titanium alloy (TNTZ) drawn wires. J Jpn Inst Met 2005;69(7):530–7.