Compressive Fatigue Behaviour of a Polymer-Based Osteochondral Scaffold

Y-H. Hsua, C Luptona, J. Tonga, A. Cosseyb, A. Auc

a Mechanical Behaviour of Materials Laboratory, School of Engineering, University of Portsmouth, UK

b Queen Alexandra Hospital, Portsmouth, UKc Smith & Nephew, Advanced Surgical Devices, Andover, MA, USA

Abstract

Cyclic loading from daily activities or from intensive exercise can lead to increased risk of

fracture. Implants designed for load bearing purposes, such as repair of articular cartilage and

underlying subchondral bone in knees must have the necessary resistance to fatigue loadings.

To date, virtually no studies have been reported on the fatigue behaviour of osteochondral

scaffolds, despite damage due to repeated loading is of considerable interest for tissue repair

purposes.

In this study, a polymer-based osteochondral scaffold has been studied under cyclic loading

conditions. Multi-step cyclic tests have been carried out with increasing compressive loads in a

phosphate buffered saline solution at 37oC. Changes in secant modulus and residual strain

accumulation are monitored. Morphological parameters of the scaffold are determined using

micro-computed tomography. The secant modulus and the number of cycles to failure of the

scaffold are obtained and compared with those of human trabecular bone (Topolinski et al,

2011).

1. Introduction

A variety of biological and synthetic scaffolds have been developed for articular cartilage and

subchondral bone repair purposes.1-11 However, scaffolds often lack the physical structure and

mechanical properties necessary to sustain long-term applications.12 Although biocompatibility

and bio-related issues have been studied widely, evaluation of mechanical properties has been

an area of underdevelopment, with testing parameters and conditions varied greatly; and

testing under controlled physiological loading conditions similar to those in vivo particularly

lacking. Evaluation of the mechanical properties of scaffolds is not always done in aqueous

solution at 37°C, although the mechanical properties of hydrated scaffold samples are known to

be much lower than those of dry samples.4-6 13 A number of scaffolds and bones were tested in

tension,8, 14, 15 which bears little resemblance to the in vivo loading environment primarily in

compression. Human bone is subjected to a variety of loading patterns during daily routine

activities. A typical loading condition is repeated or cyclic loading as experienced in walking or

running. Bone response under such loading conditions may be referred to as fatigue

behaviour.16, 17 Fatigue damage is a process of gradual mechanical degradation caused by

repeated loading at stress or strain much lower than those required to fracture in a single

application of force. As a result, damage due to cyclic loading is of significant interest for design

and application of synthetic scaffolds.

In a conventional fatigue test, specimens are subjected to a constant amplitude load till

failure to determine the number of cycle to failure at a given load range. A number of studies

based on this method have been carried out on the fatigue behaviour of bones.15, 18-30 Some

studies have also been carried out for synthetic biopolymers.31, 32 Using this method a great

number of samples may be needed to obtain the fatigue behaviour under variable stress

amplitudes. Stepwise-testing has since been developed which would allow the generation of

fatigue data with only a small numbers of samples.33, 34 In stepwise-tests, samples are subject

to loads of a range of amplitudes in each a number of cycles are allowed to elapse, thus

providing an accelerated testing regime with increased data generation. One important

parameter obtained from such tests is the evolution of secant modulus 21, 33, 35 which is an

indication of damage accumulation.30, 36

Polymer-based osteochondral scaffolds are synthetic resorbable implants designed to

replace worn-out cartilage surfaces, restoring mobility and relieving joint pain. Although used in

some clinical cases,37-39 there is a lack of published data on the biomechanical properties of the

implant. Knowledge of its mechanical behaviour is essential in order to obtain accurate

prediction of stresses and deformation in clinical applications, where the structural integrity of

the implant is particularly important when cyclic loading is considered. The motivation of the

study is to evaluate the fatigue behaviour of an osteochondral scaffold using the step-wise

testing method and to compare the results with those from human bone in terms of stress-strain

response and secant modulus. Morphological parameters of the scaffold were also obtained

and compared with those of human bone.

2. Materials and methods

2.1 Specimen

A commercial scaffold TRUFIT (Smith and Nephew) was used in the study. It is composed of

polylactide-co-glycolide (PLG) copolymer with calcium-sulfate and polyglycolide (PGA) fibres.

The dual-layer design of the implant contains both a cartilage and a bone phase. Cylinders of

the two layers scaffolds were sectioned to obtain a single layer of TRUFIT Bone, with a

diameter of 10.5mm, a length of 9 mm and an aspect ratio of 0.86. The aspect ratio was

selected based on the information obtained previously,40-42 where reported aspect ratios are

generally less than 1. 33, 43-45 The aspect ratio and dimensions of the TRUFIT Bones are in a

similar range of those of human trabecular bones in Topolinski’s study.33 The samples were

soaked in phosphate buffered saline (PBS) solution overnight prior to testing.

2.2 Morphological Parameters

The microstructure of the specimens was examined prior to fatigue testing using micro-focus

computed tomography (CT X-Ray Inspection System, Metris X-Tek Systems Ltd). The Micro-

CT scanner was operated at a voltage of 51 kV, current of 160 μA and a voxel size of 17-20

μm. Bone volume density (BV/TV), mean thickness of the trabeculae in the specimen (Tb.Th)

and the mean distance between trabeculae (Tb.Sp) were obtained by processing the micro-CT

images using Microview software.

2.3 Compression fatigue testing

Fatigue testing was carried out on a Bose ElectroForce 3200 Testing system equipped with an

environmental chamber to which PBS solution was filled and controlled at a temperature of

37˚C. Each specimen was placed carefully at the centre of the platens inside the environmental

chamber. During the testing, the specimens were unconstrained33 between the platens of the

testing machine as fixation of the specimens to the test platens may increase stiffness46 Five

TRUFIT Bones were tested under load control26, 27 with a stepwise increase of maximum stress

whilst keeping the minimum stress constant, following a testing protocol of Topolinski et al.33

A preload of 5 N was applied first to ensure good end contact and this preload was

subsequently kept constant as the minimum load throughout the test. The maximum load

started from 10 N with a gain of 10 N at successive steps, as shown in Figure 1. The maximum

load was kept constant during a period of 500 cycles and increased to the next level thereafter.

The loading scheme was adopted from Topolinski et al33 to allow easy comparison with the

results from human trabecular bone.33 The loading waveform was sinusoidal at a frequency of 1

Hz.

Engineering stress and strain were calculated by using the applied load, displacement

and the original dimension of the specimens, where lateral deformation under compression was

not considered. The changes in the secant modulus and the maximum and minimum strains

with fatigue loading were recorded. The change in the secant modulus is considered indicative

of damage accumulation in the specimen30, 36 in fatigue studies of bones from the literature.15, 18,

21, 23-29, 33-35, 47 The secant modulus at a given cycle was defined as Δσ/Δ = (σmax-σmin)/(max-min),

determined from the stress-strain curve. Figure 2 illustrates this as the slope of the line

connecting the lowest and the highest point of a stress-strain loop. The secant modulus and

strains were measured throughout the test. The cyclic stress-strain curves were also compared

with those obtained under monotonic loading conditions.

3. Results and discussion

Table 1 shows the morphometric values of the TRUFIT Bone and human trabecular

bone.33 The average BV/TV of the scaffold (0.343) is well within the range of human trabecular

bone (0.076-0.460). The trabecular thickness of the scaffold is slightly lower (0.086mm) than

that of human trabecular bone (0.105-0.268mm); whilst the trabecular spacing (0.166mm) is

smaller than that of human (0.424-1.829mm). These indicate that the microstructure of the

scaffold consists of thinner but more densely packed struts than that of human trabecular bone.

Table 1 also includes the maximum stress corresponding to the maximum secant modulus.

Samples with higher values of the maximum stress indicate higher fatigue resistance. It is clear

that the fatigue resistance of the scaffold at an average of 0.462MPa is much lower than the

range (0.859-12.28MPa) from human trabecular bone. A large variation in the results from

human bones may be due to the varied sources (e.g., multiple donors with individual

characteristics and pathologies).33 In contrast, synthetic TRUFIT bones have more consistent

morphometric values. Although a relationship between bone structure and fatigue strength for

human bone was reported,33 it is not possible to obtain such a relation for the scaffold due to

the lack of variety of the microstructure features.

Typical stress-strain curves of a sample tested under the multi-step loading scheme

(LS1 to LS8) are shown in Figure 3. Each block, shown in a different colour, represents the

response to a single loading step with 500 loading cycles. A distinctive characteristic at all load

levels is the accumulation of residual strain, or cyclic creep, with the increase of loading cycles.

Greater strain ratchetting appears to be associated with higher load levels, when non-linear

deformation becomes more evident in the larger hysteresis loops under LS7 and LS8. This is

consistent with the behaviour observed in trabecular bones.26, 29, 35, 47

The evolution of the secant modulus (as defined in Figure 2) with cycles is shown in

Figure 4(a) for the five samples tested. In all cases, it appears that an increase in secant

modulus with cycle persisted for most of the loading steps, particularly the early ones. Only

towards later at higher load levels reductions in secant modulus become evident. The initial

hardening seems to be somewhat unexpected, certainly in contrast to the continuing increase

in the residual strain with cycle, as shown in Figure 4(b), where the maximum and the minimum

compressive strains of the five samples are recorded. Interestingly samples with lower residual

strains (A, B, C) tend to have higher secant modulus (Figure 4(a)). Both maximum and

minimum strains increased with the number of cycles during the entire test and the gap

between the maximum and the minimum strains is roughly constant for each sample, indicating

that the strain range stays constant. In addition, the strain rate increases with the increase of

the number of cycles. It is known that residual strain or cyclic creep plays an important role in

material failure due to the accumulated plastic deformation.30, 35 Increase in strains with the

number of cycles have been generally observed when testing bones under cyclic loading, 27, 34

although increasing strain rate during the entire test was found in bovine trabecular bone.34

Three stages of deformation during compression fatigue tests have been reported for human

trabecular bone,27 including a transient behaviour characterised by rapid strain increase within

the initial load cycles, saturation of strain and an acceleration towards final catastrophic failure.

When secant modulus of the scaffold is compared with those of human trabecular

bones,33 as shown in Figure 4(c), much lower secant modulus and total number of cycles to

failure were obtained for the scaffold, although the patterns of evolution with cycle appear to be

similar between the two. The initial hardening of the bone is also consistent with the results

under conventional constant-amplitude fatigue testing. Linde and Hvid51 reported that in human

trabecular bones the stiffness increases until a stress level about 50% of ultimate stress is

reached followed by decreasing stiffness, although others reported decreasing in modulus from

the beginning to the end.18, 22, 24-27, 52, 53 Michel et al35 studied the fatigue behaviour of bovine

trabecular bones under load control using a sinusoidal compressive load profile at a frequency

of 2 Hz. The results showed that the pattern of modulus change with the number of cycles was

associated with the load level. At low cyclic load levels the modulus increased initially followed

by a rapid drop in the final stage. At high cyclic loads a continuous decrease in modulus from

the beginning was found. The authors35 suggested that both creep and damage accumulation

may be responsible for fatigue failure of trabecular bone, i.e. bone fails by creep under high

cyclic loads whilst by microcrack damage accumulation at low cyclic loads. A more recent study

suggested, however, that creep effects are negligible in fatigue loading cases for trabecular

bone.54 An initial increase in modulus has been observed for TRUFIT at most of the load levels

except the highest levels, which appears to be consistent with the observation in some of the

reports.35 and 51

Stress-strain loops of first cycles at the beginning of each fatigue loading step are

presented in Figure 5(a). Each centre of the loops was moved to the origin (0,0) then the points

at the top of each loop were linked to form the initial cyclic stress-strain curve. Figure 5(b)

shows the final cyclic stress-strain curve which was obtained similarly using the loops of the

final cycles at the end of each loading step. These cyclic stress-strain loops at the first and the

final cycles do not appear to differ significantly (Figure 6, CSS_1 and CSS_f), suggesting stress

softening may not be significant. Figure 6 shows a comparison of cyclic and monotonic stress-

strain curves for a typical specimen of TRUFIT. Here LS1_0 to LS5_0 are monotonic stress-

strain curves whilst CSS_1 and CSS_f are the first and the final cyclic stress-strain curves. The

monotonic curves are the loading parts of a stress-strain curve at the beginning of the each

loading step. The slopes of these curves are the modulus of elasticity. A comparison of the

mean elastic modulus of the stress-strain curves is shown in Figure 7. The elastic modulus of

monotonic stress-strain curves increase with the number of loading steps (a) and the moduli of

cyclic curves are generally higher than those of monotonic ones, suggesting stress hardening,

consistent with the trend shown in Figure 4(a). However, the elastic modulus of first cyclic

stress-strain curve is higher than that of final cyclic stress-strain curve, suggesting damage

accumulation due to fatigue.

4. Conclusions

Multi-step cyclic tests have been carried out on TRUFIT Bones to study the fatigue behaviour of

the scaffold under increasing compressive cyclic loading conditions. Morphological parameters

of TRUFIT scaffold have also been obtained using computed tomography and compared with

those of human trabecular bone. The results show an increase in secant modulus and stress

hardening during the initial steps of fatigue loading, consistent with that observed in human

trabecular bone. The modulus and the number of cycle to failure of the scaffold are, however,

significantly lower than those of human trabecular bone. The elastic modulus of monotonic

stress-strain curves increases with the number of loading steps and the moduli of cyclic curves

are generally higher than those of monotonic curves. Progressive increase in residual strains

has been observed in the scaffold for the entire test duration, indicating that residual strain

accumulation, or cyclic creep, may be the predominant driving force leading to failure of the

scaffold.

Multi-step tests seem to be useful in assessing the essential fatigue behaviour of

scaffolds. These tests allow the reduced number of samples for a reliable analysis of fatigue

properties of biomaterials.

Acknowledgements

The authors would like to thank Smith & Nephew for the provision of samples and University of

Portsmouth for financial support.

References1. Toolan BC, Frenkel SR, Pachence JM, Yalowitz L, Alexander H. Effects of growth-factor-

enhanced culture on a chondrocyte-collagon implant for cartilage repair. Journal of Biomedical

Materials Research 1996;31:273-80.

2. Athanasiou K, Schmitz JP, Agrawal CM. The effects of porosity on in virto degradation of

polylactic-acid-polyglycolic acid implants used in repair of articular cartilage. Tissue

Engineering 1998;4(1):53-63.

3. Lu L, Zhu X, Valenzuela RG, Currier BL, Yaszemski MJ. Biodegradable polymer scaffolds for

cartilage tissue engineering. Clinical Orthopaedics and Related Research 2001;391s:S251-

S70.

4. Slivka MA, Leatherbury NC, Kieswetter K, Niederauer GG. Porous, resorbable, fiber-

reinforced scaffolds tailored for articular cartilage repair. Tissue Engineering 2001;7(6):767-80.

5. Harley BA, Lynn AK, Wissner-Gross Z, Bonfield W, Yannas IV, Gibson LJ. Design of a

multiphase osteochondral scaffold. II. Fabrication of a mineralized collagen–glycosaminoglycan

scaffold. Journal of Biomedical Materials Research Part A 2010;92(3):1066-77.

6. Harley BA, Lynn AK, Wissner-Gross Z, Bonfield W, Yannas IV, Gibson LJ. Design of a

multiphase osteochondral scaffold III: Fabrication of layered scaffolds with continuous

interfaces. Journal of Biomedical Materials Research Part A 2010;92(3):1078-93.

7. Oka M, Ushio K, Kumar P, Ikeuchi K, Hyon SH, Nakamura T et al. Development of artificial

articular cartilage. Proceedings of the Institution of Mechanical Engineers, Part H: Journal of

Engineering in Medicine 2000;214:59-68.

8. Bera B. Development of artificial articular cartilage. S¯adhan¯a. 2009;34(5):823-31.

9. Oka M, Noguchi T, Kumar P, Ikeuchi K, Yamamuro T, Hyon SH et al. Development of an

artificial articular cartilage. Clinical Materials 1990;6:361-81.

10. Hung CT, Lima EG, Mauck RL, Taki E, LeRoux MA, Lu HH et al. Anatomically shaped

osteochondral constructs for articular cartilage repair. Journal of Biomechanics

2003;36(12):1853-64.

11. Sherwood JK, Riley SL, Palazzolo R, Brown SC, Monkhouse DC, Coates M et al. A three-

dimensional osteochondral composite scaffold for articular cartilage repair. Biomaterials

2002;23(24):4739-51.

12. Suh J-K, Arøen A, Muzzonigro TS, Disilvestro M, Fu FH. Injury and repair of articular

cartilage: Related scientific issues. Operative Techniques in Orthopaedics 1997;7(4):270-8.

13. Slivka MA, Leatherbury NC, Kieswetter K, Niederauer GG. In vitro compression testing of

fiber-reinforced, bioabsorbable, porous implants. In: Mauli Agrawal C, Parr JE, Lin ST, editors.

Synthetic Bioabsorbable Polymers for Implants. Kansas City, USA2000. p. 124-35.

14. Messner K. Hydroxylapatite supported Dacron plugs for repair of isolated full-thickness

osteochondral defects of the rabbit femoral condyle: mechanical and histological evaluations

from 6-48 weeks. Journal of Biomedical Materials Research 1993;12:1527-32.

15. Zioupos P, Casinos A. Cumulative damage and the response of human bone in two-step

loading fatigue. Journal of Biomechanics 1998;31:825-33.

16. Martin RB. Fatigue microdamage as an essential element of bone mechanics and biology.

Calcified Tissue International 2003;73:101-7.

17. Taylor M, Tanner KE. Fatigue failure of cancellous bone: a possible cause of implant

migration and loosening. The Journal of Bone & Joint Surgery. 1997;79-B:181–2.

18. Pattin CA, Calert WE, Carter DR. Cyclic mechanical property degradation during fatigue

loading of cortical bone. Journal of Biomechanics 1996;29(1):69-79.

19. O’Brien FJ, Taylor D, Lee TC. Microcrack accumulation at different intervals during fatigue

testing of compact bone. Journal of Biomechanics. 2003;36:973-80.

20. Yeni YN, Fyhrie DP. Fatigue damage-fracture mechanics interaction in cortical bone. Bone.

2002;30(3):509-14.

21. Cotton JR, Winwood K, Zioupos P, Taylor M. Damage rate is a predictor of fatigue life and

creep strain rate in tensile fatigue of human cortical bone samples. Journal of Biomechanical

Engineering 2005;127:213-9.

22. Fleck C, Eifler D. Deformation behaviour and damage accumulation of cortical bone

specimens from the equine tibia under cyclic loading. Journal of Biomechanics 2003;36:179-89.

23. Yamamoto E, Crawford RP, Chan DD, Keaveny TM. Development of residual strains in

human vertebral trabecular bone after prolonged static and cyclic loading at low load levels.

Journal of Biomechanics 2006;39:1812-8.

24. Zioupos P, Wang XT, Currey JD. The accumulation of fatigue microdamage in human

cortical bone of two different ages in vitro. Clinical Biomechanics. 1996;11:365-75.

25. Moore TLA, Gibson LJ. Fatigue microdamage in bovine trabecular bone. Journal of

Biomechanical Engineering 2003;125:769-76.

26. Moore TLA, Gibson LJ. Fatigue of bovine trabecular bone. Journal of Biomechanical

Engineering 2003;125:761-8.

27. Dendorfer S, Maier HJ, Taylor D, Hammer J. Anisotropy of the fatigue behaviour of

cancellous bone. Journal of Biomechanics 2008;41:636-41.

28. Carter DR, Hayes WC. Compact bone fatigue damage – I. Residual strength and stiffness.

Journal of Biomechanics 1977;10:325-37.

29. Rapillard l, Charlebois M, Zysset P. Compressive fatigue behavior of human vertebral

trabecular bone. Journal of Biomechanics 2006;39:2133-9.

30. Dendorfer S, Maier HJ, Hammer J. Fatigue damage in cancellous bone: an experimental

approach from continuum to micro scale. Journal of the Mechanical Behavior of Biomedical

Materials 2009;2(1):113-9.

31. Palissery V, Taylor M, Browne M. Fatigue characterization of a polymer foam to use as a

cancellous bone analog material in the assessment of orthopaedic devices Journal of Materials

Science: Materials in Medicine 2004;15(1):61-7.

32. Meyer RW, Pruitt LA. The effect of cyclic true strain on the morphology, structure, and

relaxation behavior of ultra high molecular weight polyethylene Polymer 2001;42:5293-306.

33. Topolinski T, Cichanski A, Mazurkiewicz A, Nowicki K. Study of the behavior of the

trabecular bone under cyclic compression with stepwise increasing amplitude. Journal of the

Mechanical Behavior of Biomedical Materials 2011;4(8):1755-63.

34. Guillén T, Ohrndorf A, Tozzi G, Tong J, Christ H-J. Compressive fatigue behavior of bovine

cancellous bone and bone analogous materials under multi-step loading conditions. Advanced

Engineering Materials 2012;14(5):B199-B207.

35. Michel MC, Guo X-DE, Gibson LJ, McMahon TA, Hayes WC. Compressive fatigue behavior

of bovine trabecular bone. Journal of Biomechanics 1993;26(4-5):453-63.

36. Zioupos P, Gresle M, Winwood K. Fatigue strength of human cortical bone: Age, physical,

and material heterogeneity effects. Journal of Biomedical Materials Research Part A

2008;86A(3):627-36.

37. Carmont MR, Carey-Smith R, Saithna A, Dhillon M, Thompson P, Spalding T. Delayed

incorporation of a Trufit Plug: perseverance is recommended. Arthroscopy: The Journal of

Arthroscopic and Related Surgery 2009;25(7):810-4.

38. Dhollander AAM, Liekens K, Almqvist KF, Verdonk R, Lambrecht S, Elewaut D et al. A pilot

study of the use of an osteochondral scaffold plug for cartilage repair in the knee and how to

deal with early clinical failures. Arthroscopy: The Journal of Arthroscopic and Related Surgery

2012;28(2):225-33.

39. Pearce CJ, Gartner LE, Mitchell A, Calder JD. Synthetic osteochondral grafting of ankle

osteochondral lesions. Foot and Ankle Surgery 2012;18:114-8.

40. DiSilvestro MR, Suh J-KF. A cross-validation of the biphasic poroviscoelastic model of

articular cartilage in unconfined compression, indentation, and confined compression. Journal

of Biomechanics 2001;34:519-25.

41. Huang C-Y, Stankiewicz A, Ateshian GA, Mow VC. Anisotropy, inhomogeneity, andtension–

compression nonlinearity of human glenohumeral cartilage in finite deformation. Journal of

Biomechanics 2005;38:799-809.

42. Linde F, Nobrgaard P, Hvid I, Odgaard A, Soballe K. Mechanical properties of trabecular

bone. Dependency on strain rate. Journal of Biomechanics 1991;24(9):803-9.

43. Carter DR, Hayes WC. The compressive behavior of bone as a two-phase porous structure.

The Journal of Bone and Joint Surgery (Am) 1977;59(7):954-62.

44. Korhonen RK, Laasanena MS, Toyras J, Rieppob J, Hirvonena J, Helminen HJ et al.

Comparison of the equilibrium response of articular cartilage in unconfined compression,

confined compression and indentation. Journal of Biomechanics 2002;35:903-9.

45. Park S, Hung CT, Ateshian GA. Mechanical response of bovine articular cartilage under

dynamic unconfined compression loading at physiological stress levels. OsteoArthritis and

Cartilage 2004;12:65-73.

46. Linde F, Hvid I. The effect of constraint on the mechanical behaviour of trabecular bone

specimens. Journal of Biomechanics 1989;22(5):485-90.

47. Haddock SM, Yeh OC, Mummaneni PV, Rosenberg WS, Keaveny TM. Similarity in the

fatigue behavior of trabecular bone across site and species. Journal of Biomechanics

2004;37:181-7.

48. Caler WE, Carter DR. Bone creep-fatigue damage accumulation. Journal of Biomechanics

1989;22:625-35.

49. Topolinski T, Cichanski A, Mazurkiewicz A, Nowicki K. Study of the behavior of the

trabecular bone under cyclic compression with stepwise increasing amplitude. Journal of the

Mechanical Behavior of Biomedical Materials In Press.

50. Fray ME, Altstadt V. Fatigue behaviour of multiblock thermoplastic elastomers. 1. Stepwise

increasing load testing of poly(aliphatic/aromatic-ester) copolymers. Polymer. 2003;44:4635-42.

51. Linde F, Hvid I. Stiffness behaviour of trabecular bone specimens. Journal of Biomechanics

1987;20(1):83-9.

52. Schaffler MB, Radin EL, Burr DB. Mechanical and morphological effects of strain rate on

fatigue of compact bone. Bone 1989;10:207-14.

53. Zioupos P, Wang XT, Currey JD. Experimental and theoretical quantification of the

development of damage in fatigue tests of bone and antler. Journal of Biomechanics

1996;29:989-1002.

54. Moore TLA, O’Brien FJ, Gibson LJ. Creep does not contribute to fatigue in bovine

trabecular bone. Journal of Biomechanical Engineering 2004;126(3):321-9.

Figure1. The loading scheme of a stepwise fatigue test (Fmax=10N; Fmin=Fpreload=5N), following the protocol of Topolinski et al, 2011.33

Figure 2. Secant modulus (Ei) is defined as the stress range divided by the strain range at a given cycle i, which is determined by the slope of the line connecting the minimum and the maximum stress values of the cycle.

Figure 3. Stress-strain curves from a multi-step fatigue test.

Figure 4. (a) The evolution of secant modulus as a function of cycles for TRUFIT Bone (b) the development of

the maximum and minimum strains with cycle and (c) comparison of secant modulus of TRUFIT Bone and human trabecular bone.33

Figure 5. Cyclic stress-strain responses: (a) Loops of first cycles at the beginning of each loading step; (b) loops of the final cycles at the end of each loading step.

Figure 6. Comparison of the cyclic and the monotonic stress-strain curves of a typical sample.

Figure 7. The elastic modulus of (a) monotonic curves; and (b) cyclic stress-strain curves.

Table 1. Morphometric values and maximum stresses of TRUFIT bone and human trabecular bone. 33

TRUFIT BoneBone volume density

BV/TVTrabecular thickness

Tb.Th (mm)Trabecular spacing

Tb.Sp (mm)

Max Stress (MPa)(corresponding to

Emax)

A 0.359 0.089 0.161 0.547

B 0.340 0.082 0.158 0.328

C 0.353 0.087 0.159 0.547

D 0.328 0.089 0.184 0.337

E 0.333 0.084 0.169 0.552

Human trabecular bone [33]

0.076 - 0.460 0.105 - 0.268 0.424 - 1.829 0.859 - 12.280