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Tensile damage and its effects on cortical bone
ARTICLE IN PRESS
Journal of Biomechanics 36 (2003) 1683–1689
Tensile damage and its effects on cortical bone S.P. Kothaa, N. Guzelsub,* b
a Biomechanics and Biomaterials Laboratory, University of Missouri—Kansas City, 650 E. 25th Street, Kansas City, MO 64108, USA Biomechanics Laboratory, University of Medicine and Dentistry of New Jersey, SOM, Tr. #4, 675 Hoes Lane, Piscataway, NJ 08854, USA
Accepted 2 April 2003
Abstract Plexiform bovine bone samples are repeatedly loaded in tension along their longitudinal axis. In order to induce damage in the bone tissue, bone samples are loaded past their yield point. Half of the bone samples from the damaged group were stored in saline to allow for viscoelastic recovery while the others were decalcified. Tensile tests were conducted on these samples to characterize the effects of damage on the mechanical behavior of the organic matrix (decalcified samples) as well as on bone tissue (stored in saline). The ultimate strain of the damaged decalcified bone is 29% higher compared to that of non-damaged decalcified (control) bone. The ultimate stresses as well as the elastic moduli are similar in both decalcified groups. This phenomenon is also observed in other collagenous tissue (tendon and ligament). This may suggest that damage in bone is caused by shear failure of the organic matrix; transverse separation of the collagen molecules or microfibrils from each other. In contrast, there is a trend towards lowered ultimate strains in damaged bone, which is soaked in saline, with respect to control bone samples (not damaged). The damaged bone tissue exhibits a bi-linear behavior in contrast to the mechanical behavior of non-damaged bone. The initial elastic modulus (below 55 MPa) and ultimate strength of damaged bone are similar to that in non-damaged bone. r 2003 Elsevier Ltd. All rights reserved. Keywords: Plexiform bone; Decalcified bone; Shear failure; Damage in type I collagen; Organic matrix
1. Introduction Bone microcracks are most likely manifestations of fatigue damage that develop under repetitive loading conditions (Carter and Hayes, 1977a, b; Schaffler et al., 1990, 1995; Wenzel et al., 1996; Zioupos et al., 1996; Pattin et al., 1996; Boyce et al., 1998; Burr et al., 1998; Zioupos and Casinos, 1998; Fazzalari et al., 1998; Donahue et al., 2000; Yeni and Fyhrie, 2002). Numerous microcracks (o10 mm in length) disperse throughout the bone specimen when they are loaded in tension beyond their yield point (Zioupos et al., 1994; Zioupos and Currey, 1994) along with some larger cracks (Burr et al., 1998; O’Brien et al., 2000). There are no studies that illustrate the nanostructural evolution of damage. Different scenarios of damage initiation leading to *Corresponding author. Department of Osteosciences/Biomechanics, Unviersity of Medicine and Dentistry, PCC Suite 102, 40 East Llaurel Road, Stratford, NJ 08084-1504, USA. Tel.: +1-856-566-2731; fax: +1-856-566-2733. E-mail address: [email protected] (N. Guzelsu). 0021-9290/03/$ - see front matter r 2003 Elsevier Ltd. All rights reserved. doi:10.1016/S0021-9290(03)00169-6
cracks in bone tissue have been suggested (Nicolella et al., 1997; Yeni and Fyhrie, 2002). Damage in the form of cracks can start and develop by tensile failure of collagen molecules/fibrils and/or mineral, by shear failure between collagen molecules/fibrils, and/or by debonding of the organic matrix from the mineral platelets (Nicolella et al., 1997; Yeni and Fyhrie, 2002). A better understanding of the nanoscale damage mechanism in bones due to excess repetitive loads may help physicians to devise treatments for stress fractures more effectively in athletes and army personnel (Macera, 1992; Jones et al., 1993; Ross, 1993; Brill and Macera, 1995; Schaffer et al., 1999; Beck et al., 2000; Givon et al., 2000; Popovich et al., 2000; Marx et al., 2001) as well as in the prediction of fracture risk in osteoporotic bones. Studies of the distribution of stresses and strains within the organic and mineral components may illuminate the nanostructural origin of damage in bone. Cortical bone is considered a short fiber reinforced composite where hydroxyapatite platelets reinforce the organic matrix (primarily type I collagen) (Currey, 1964,
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1969a, b; Carter and Spengler, 1978). When bone is subjected to axial tensile stresses, the organic matrix is subjected to both normal and shear stresses (Currey, 1964, 1969a, b; Carter and Spengler, 1978; Wagner and Weiner, 1992; Jager and Fratzl, 2000; Kotha et al., 2000; Kotha and Guzelsu, 2002) that may cause nanoscale damage in the organic matrix of bone. Further delving into the hierarchical organization of the organic matrix indicates that the organic matrix, which is made up of mainly type I collagen, acts as a composite material itself. The collagen fibrils transfer the tensile forces from one fibril to the next through shear stresses that are generated along the overlap region of the fibrils (Birk et al., 1991; Silver, 1987). Deformation of collagen fibers involves molecular stretching and slippage, fibrillar slippage, and, ultimately, defibrillation (Pins et al., 1997). These mechanisms decrease the fibril diameter and increase the toe region during subsequent tensile testing. Studies of the nanoscale damage mechanism in connective tissues made of type I collagen (tendons or ligaments) have been conducted by mechanically loading these tissues and introducing damage in their ultrastructure (Kastelic and Baer, 1980; Provenzano et al., 2002). Damaged tendons (damaged by loading the tendon up to a certain point and releasing the load) have increased strains to failure compared to nondamaged tendons (Kastelic and Baer, 1980). Provenzano and his co-workers have shown that damaged ligaments (damaged by pulling to a threshold value of 5.14% strain and allowed to recover) demonstrate an increased low-strain region deformation compared to intact samples (increased toe region) (Provenzano et al., 2002). Subfailure damage has been shown to alter the mechanical properties of the anterior cruciate ligament and lengthen the toe region of the force–displacement curve, which result in increased joint laxity (Panjabi et al., 1996). These studies support Kastelic and Baer’s observation that nanoscale damage in the collagen structure increases crimps and results in increased toe region deformation (Kastelic and Baer, 1980). We investigate if these phenomena are also observed in the organic matrix of damaged bone tissue. In this study, we introduce a new method to assay the nanoscale damage phenomena in the organic matrix of bone tissue based on the above studies. We induce tensile damage in cortical bone by repetitively loading it past its yield point. We test the hypothesis that damaged bone tissues exhibit a bi-linear stress–strain behavior while non-damaged bone tissues exhibit a linear stress– strain behavior before their yield point, when calcified samples from these two groups are tested in tension. Furthermore, we test the hypothesis that the ultimate strain of the organic matrix of damaged bone tissue is greater than that of the organic matrix of non-damaged bone tissue when decalcified samples from the two groups are tested in tension.
2. Methods Two-year-old fresh cow femur diaphysis, which were obtained from a local slaughterhouse (Moyer Packing, Sauderton, PA) were used to make tensile mechanical test specimens. The plexiform nature of the samples was confirmed through optical microscopy. Mechanical tests were performed on an Instron 1321B universal testing machine and data was recorded at 60 Hz. The range of the load cell for calcified specimens was 2000 N (resolution of 2.0 N) while for testing decalcified specimens the range was 1000 N (resolution of 1.0 N). In these experiments, 40 samples were divided into four groups. Dumbbell shape mechanical test specimens were made from the lateral and medial cortices of femurs according to ASTM standards (ASTM, 1998; DePaula et al., 2002; with a 2.0 mm 5.0 mm cross-section of the gauge length). Two experimental group samples were repeatedly loaded (eight cycles) in tension past their yield point under stroke control (at a strain rate of 0.0005 s 1) using a triangular waveform to induce damage. The yield point was determined using the 0.2% offset strain method. While they were repeatedly loaded, the average maximum stresses and strains in the bone samples were 110 MPa (110713.8 Mpa–Mean7SD) and 0.79% (0.7970.12%), respectively. Experiments were conducted carefully to prevent the samples from being loaded in compression. Two samples broke during this procedure resulting in nine samples in each of the two damaged experimental groups. The strain in the mineralized samples was measured using an extensometer with a gage length of 12.5 mm. The range and resolution of the extensometer were set to 70.5 and 0.0005 mm, respectively. The recorded stress–strain curves confirmed that the specimens were loaded past their yield point (Fig. 1). Nine damaged and 10 non-damaged samples were decalcified in 0.5 M EDTA solutions for 10 days. Solutions were changed every 2 days (pH=8.3 at 4 C). Each specimen was decalcified in 150 ml solution (Catanese et al., 1999). Sample ends were coated with paraffin to expose a 14 mm rectangular region in the middle of the samples. With this technique, the ends of the samples were kept intact. Decalcification was confirmed as only 0.2% by weight of calcium was eluted in 1 N hydrochloric acid solution as measured by absorptiometry (Sarkar and Chauhan, 1967). These samples were equilibrated for 1 day at 37 C in physiological saline (0.145 M NaCl, pH=7.5) before being mechanically tested. The remaining damaged nine samples and non-damaged intact specimens (10 of them) were immersed in 0.145 M saline at pH=7.5 and 37 C for 8 days. Immersion of damaged bones in saline solution for 8 days allowed the bone samples to recover their viscoelastic deformation after repeated loading
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beyond their yield point. Immersion of bones in saline for a long period of time causes a small amount of bone mineral dissolution, leading to a small reduction in the mechanical properties (Gustafson et al., 1996). Dissolution of bone mineral does not cause a bi-linear stress– strain curve type behavior (DePaula, 2000). In our tension experiment, we compared the mechanical behavior of damaged (experimental calcified group—to allow for viscoelastic recovery) and non-damaged (calcified control—similar equilibration time as experimental group) bones after immersion in saline for 8 days. All four groups were tested in tension under stroke control at a strain rate of 0.0005 s 1. The samples were
Fig. 1. Stress–strain behavior of the same sample tested at a strain rate of 0.0005 s 1. The average maximum stresses and strains in the bones while they are repeatedly loaded are 110.4713.8 MPa and 0.7970.12%, respectively. Note the bi-linear stress–strain behavior of damaged bone after immersion in saline for a period of 8 days. The slope of the upper part of the stress–strain curve of damaged bone is significantly lower at the end of the eighth cycle (Po0:005) due to repeated tensile loading. All values are mean7SD.
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kept wet during the testing and they were loaded to failure in tension. Strains in the decalcified samples were obtained from the increase in displacement between the grips for samples with gage lengths of 12.5 mm (range and resolution of the displacement transducer were 75 and 0.005 mm, respectively). For statistical analysis, one-tailed t-tests were performed to obtain significance.
3. Results The damaged calcified bones had a bi-linear stress– strain behavior before the macroscopic yield point (Fig. 1). These specimens showed a slope transition at about 50–60 MPa. Bi-linear stress–strain behavior is not clearly observable for the 8th cycle of initial loading (Fig. 1, Table 1). Moreover, there was a trend for the damaged bones to exhibit lower ultimate strains (by B30%) in comparison with non-damaged specimens (Fig. 2, Table 1—Columns 5 and 6, Po0:055; post hoc analysis indicated power=0.53 at alpha=0.05). There was a trend towards lower toughness in damaged specimens compared to non-damaged specimens because of lower ultimate strains in damaged specimens and similar ultimate stresses of both groups (Po0:07; post hoc analysis indicated power=0.59 at alpha=0.05). The initial elastic modulus of the control samples and the damaged specimens were similar (Table 1—Column 3). On the other hand, above the 55 MPa stress level, the elastic modulus of the damaged samples was 40% lower than the control samples (Table 1— Column 4, Po0:05; post hoc analysis indicated power=0.996 for alpha=0.05). The damaged specimens that were decalcified showed 29% higher ultimate strains compared with nondamaged decalcified specimens (Po0:05; post hoc analysis indicated power=1.000 for alpha=0.05) (Fig. 3). There were no significant differences in the
Table 1 Elastic moduli, ultimate stress and ultimate strain for the calcified bone samples Groups
Treatments and applied damaged load cycles
Initial elastic modulus below 55 MPa stress level in GPa
Elastic modulus above the 55 MPs stress level in GPa
Ultimate stress in MPa
Ultimate strain (%)
Calcified control samples
Soaked in saline for 8 days
19.673.6
17.172.9a
103.3720.9
2.5770.90c
Calcified damaged samples
First cycle
21.771.6
19.072.5b
Eighth cycle Soaked in saline for 8 days after the eighth cycle
18.372.2 20.273.1
12.472.7b 10.472.4a
106.2716.7
1.7971.01c
Similar alphabets indicate that the values are statistically different from one another (Po0:05).
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(control) versus 0.3270.12 GPa (damaged)} of the 2 decalcified groups (Fig. 3). This indicates that the increased strain in the organic matrix of damageddecalcified bones compared to control-decalcified bones occurs primarily in the toe region.
4. Discussion
Fig. 2. The stress–strain curves for control and damaged bone after being immersed in saline for 8 days. Note the bi-linear stress–strain curve for damaged bone that is not present in control bone. The slopes of the lower part of the two groups are statistically similar (both groups were stored in saline for 8 days). The slope of the upper part of the stress–strain curve of damaged bone is significantly lower than controls (Po0:001). The ultimate strains of damaged bone are significantly different from non-damaged controls (Po0:055) while the ultimate stresses are statistically similar. All values are mean7SD.
Fig. 3. Representative stress–strain curves for decalcified control and decalcified damaged bone. Note the increased ultimate strains in the decalcified organic of the damaged bone as compared with controls (Po0:005). The ultimate strengths of the damaged group are statistically similar to controls. The slopes of the stress–strain curves for both groups are statistically similar in the linear portions. It is noted that three of the samples (two in controls and one in the damaged group) had slight decrements in stresses past the maximum stress. All values are mean7SD.
ultimate strengths {19.3274.25 MPa (control) versus 21.4073.89 MPa (damaged)} or the elastic modulus between 30% and 60% level {0.3270.11 GPa (control) versus 0.2970.13 GPa (damaged)}, and the elastic modulus between 60% and 90% level {0.3670.10 GPa
The presence of damage in bone tissue has been shown to increase the ultimate strain of the organic matrix of the damaged bone as compared to the organic matrix of control bone (Fig. 3). The increase in ultimate strains of the organic matrix in damaged bone was primarily due to changes in the toe region of the stress– strain curve. In collagenous tissue (tendon), the increase in ultimate strains is attributed to a straightening of the crimps as observed with a polarizing microscope (Kastelic and Baer, 1980). We were not able to observe the increased crimps through a polarizing microscope in our decalcified samples. This may be because the organic matrix in the bone tissue is highly disoriented in comparison with the organic structure in tendon. The organic matrix of bone tissue acts to transmit forces, transfer forces among the bone mineral platelets, and prevent bone tissue from failing prematurely as a brittle material. Under excess load, deformation of collagen fibers involves stretching, slippage of laterally adjoining elements and separation (in collagen molecule and/or in collagen fibril levels) and ultimately defibrillation of the fibrils from the overlap regions under shear force transmission (Pins et al., 1997; Sasaki and Odajima, 1996; Sasaki et al., 1999; Christiasen et al., 2000). Provenzano and his co-workers (Provenzano et al., 2002) showed that the stress–strain curve of rat medial collateral ligaments (MCL) that were allowed to recover their viscoelasticity after being loaded above the damage threshold of 5.14%, exhibit an elongated toe region as well as decreased tangential modulus and ultimate stress. Strains of up to 5.14% only increased the elongation in the toe region of MCLs without affecting the elastic modulus or ultimate strengths. Kastelic and Baer also observed a similar behavior in their experiments on rat tail tendons (Kastelic and Baer, 1980). These findings support our findings, wherein, load-induced slippage or separation of some of the laterally joined microfibril structures could reduce the effective diameter of the fibrils and increase the toe region deformation of the collagen structure. In their studies (Provenzano et al., 2002), above 5.14% strains, preloading conditions might be causing lateral defibrillation due to shear forces and also the breakage in the linearly attached fibrils and/or breakage of the fibrils due to axial tension forces, which can alter the elastic modulus and ultimate stress. This behavior seems to be a characteristic of type I collagen as experiments on self
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assembled type I collagen fibrils indicate that a decrease in the diameter of the collagen microfibrils (thinner microfibrils) increases the toe region (low-strain elastic modulus) while an increase in the length of the collagen microfibrils increases the elastic modulus at higher strains as well as the ultimate strength (Christiasen et al., 2000). In a separate study (Yahia et al., 1990), rabbit MCLs subjected to different predetermined strain levels, showed that strains of 10% and above result in numerous broken thin collagen fibers and thick collagen fiber bundles under scanning electron microscopy. Below 10% strain, they did not observe disruption of the collageneous entities. This also indicates that up to certain threshold strain level, lateral separation (defibrillation) could take place in the collagen structure without breakage in linearly attached fibers (breakage across the fibrillar structure) (Christiasen et al., 2000; Provenzano et al., 2002). All these observations suggest that increased crimps are present in the decalcified organic of damaged bones due to decreased lateral association between collagen molecules/microfibrils resulting from a shear failure between collagen molecules/microfibrils as is observed in other collageneous tissues (tendon, ligaments) (Kastelic and Baer, 1980; Provenzano et al., 2002; Christiasen et al., 2000). The dissociation between adjacent microfibrils due to shear damage allows the damaged organic bone matrix to be more crimpled as in damaged tendon (Kastelic and Baer, 1980). This results in increased toe region strains during subsequent mechanical testing. One of the limitations of this study is that the amounts of denatured collagen or the cross-link content were not measured. Our measurements of the mechanical properties of non-damaged decalcified bovine bone are lower than reported by Bowman et al. (1996) for decalcified bovine bone but are within the values reported by Catanese et al. (1999) and Wang et al. (2002) for decalcified human bone tissue. The reason for this discrepancy with the results of Bowman et al. (1996) may be due to their use of preload cycles (up to 5% strain) or the age (not specified) and the location (humerus) of samples. Another limitation is that the number and size of microcracks were not measured in the non-decalcified bones. According to Reilly and Currey (1999), microcracks (o10 mm) were initiated at 0.4% tensile strain in equine bone and there was a considerable growth in microcrack density when tensile strain reached 0.8%. As the strains reached 0.78% in our loading protocol, it was believed that a large number of microcracks (o10 mm) would be present in the bovine bone used in these experiments. The stiffness of bone decreased by 35% (19.4–12.4 GPa) during loading (from first cycle to the eighth cycle) (Fig. 1) and Burr and his co-workers (Burr et al., 1998) have shown larger microcracks after a 15%
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stiffness degradation during bending of a canine femur, a smaller number of larger microcracks may also be present in our damaged samples. The bi-linear behavior of the initial part of the stress– strain curve after the first cycle (different slopes below and above 55 MPa; Fig. 1) are also observed by other researchers (Fondrk et al., 1988; Bonfield and Li, 1967). Fondrk and his co-workers suggested that that the bilinear stress–strain behavior was due to plastic flow and/ or slip behavior or incomplete crack closure (Fondrk et al., 1999). This point of view was supported by the increase in residual volume of bovine and human cortical bone loaded at high-strain amplitudes (Fondrk et al., 1999) and the presence of inelastic strains (Fondrk et al., 1988; Bonfield and Li, 1967). The nanostructural mechanisms responsible for the plastic flow, slip behavior or incomplete crack closure may be related to the damage mechanism within the collagen microfibril such that the presence of mineral within the sheared collagen microfibrils may hold it in an extended configuration, preventing it from reorganization. The mechanical behavior of non-decalcified damaged bone in this study is similar to that described in the literature for bone repeatedly loaded in tension (Zioupos et al., 1994; Jepsen et al., 2001). This behavior is better delineated after the viscoelastic strain is recovered after immersion in saline (8 days in our experiments) (Jepsen et al., 2001). Essentially, the initial elastic modulus of damaged bone is the same as that of non-damaged bone and stress–strain curve becomes curvilinear at higher stresses (>55 MPa in our case). Our results contrast with the results obtained by Carter and Hayes (1977b) in which cortical bone subjected to cyclic bending had lowered residual strength and initial elastic modulus, but agree with studies by Martin et al. (1997) where the residual strength and the initial elastic modulus did not decrease. This may be due to differences in the extent of damage induced in the bone tissue. The technique introduced in this study (decalcifying bone samples after inducing damage), allows us to investigate what happens in the organic bone matrix due to damage in bone tissue. Although we did not directly measure the micro- or macrocracks in our samples, our experimental protocol was similar to that used by other researchers to generate damage in cortical bone. Bone as a composite material is affected by the mechanical properties of both phases. The changes taking place in the organic matrix, which can be investigated with this technique, could be one of the reasons for microfracture initiation in bone samples. As suggested by Nicolella and his co-workers and, Yeni and Fyhrie; shear failure between the collagen molecules/ fibrils, supported by this study, may be one of the main factors starting microcracks in the bone tissue due to excess loading (Nicolella et al., 1997; Yeni and Fyhrie, 2002).
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ARTICLE IN PRESS S.P. Kotha, N. Guzelsu / Journal of Biomechanics 36 (2003) 1683–1689 Popovich, R.M., Gardner, J.W., Potter, R., Knapick, J.J., Jones, B.H., 2000. Effect of rest from running on overuse injuries in army basic training. American Journal of Preventive Medicine 18 (3S), 147–155. Provenzano, P.P., Heisey, D., Hayashi, K., Lakes, R., Vanderby, R., 2002. Subfailure damage in ligament: a structural and cellular evaluation. Journal of Applied Physiology 92 (1), 362–371. Reilly, G.C., Currey, J.D., 1999. The development of microcracking and failure in bone depends on the loading mode to which it is adapted. Journal of Biomechanics 202 (5), 543–552. Ross, J., 1993. A review of lower limb overuse injuries during basic military training. Part 1: types of overuse injuries. Military Medicine 158, 410–420. Sarkar, B.C.R., Chauhan, U.P.S., 1967. A new method for determining micro-quantities of calcium in biological materials. Annals of Biochemistry 20, 155–166. Sasaki, N., Odajima, S., 1996. Elongation mechanism of collagen fibrils and force–strain relations of tendon at each level of structural hierarchy. Journal of Biomechanics 29, 1131–1136. Sasaki, N., Shukunami, N., Matsushima, N., Izumi, Y., 1999. Time-resoved X-ray diffraction from tendon collagen during creep using synchroton radiation. Journal of Biomechanics 32, 285–292. Schaffer, R.A., Brodine S, K., Almeida, S.A., William, K.M., Ronaghy, S., 1999. Use of simple measures of physical activity to predict stress fractures in young men undergoing a rigorous physical training program. American Journal of Epidemiology 149, 236–242. Schaffler, M.B., Choi, K., Milgrom, C., 1995. Aging and microdamage accumulation in human compact bone. Bone 17, 521–525.
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Schaffler, M.B., Radin, E.L., Burr, D.B., 1990. Long-term fatigue behavior of compact bone at low strain magnitude and rate. Bone 11, 321–326. Silver, F.H., 1987. Biological Materials: Structure, Mechanical Properties and Modeling of Soft Tissues. New York University Press, New York, pp. 91–110. Wagner, H.D., Weiner, S., 1992. On the relationshiop between the micrstructure of bone and its mechanical stiffness. Journal of Biomechanics 25 (11), 1311–1320. Wang, X., Shen, X., Li, X., Agrawal, C.M., 2002. Age related changes in the collagen network and toughness of bone. Bone 31 (1), 1–7. Wenzel, T.E., Schaffler, M.B., Fyhrie, D.P., 1996. In vivo trabecular microcracks in human vertebral bone. Bone 19, 89–95. Yahia, L., Brunet, J., Labelle, S., Rivard, C-E., 1990. A scanning electron microscopic studt of rabbit ligaments under strain. Matrix 10, 58–64. Yeni, Y.N., Fyhrie, D.P., 2002. Fatigue damage–fracture mechanism interactions in cortical bone. Bone 30 (3), 509–514. Zioupos, P., Casinos, A., 1998. Cumulative damage and the response of human bone in two-step loading fatigue. Journal of Biomechanics 31, 825–833. Zioupos, P., Currey, J.D., 1994. Extent of microcracking and morphology of microcracks in damaged bone. Journal of Material Science 29 (4), 978–986. Zioupos, P., Currey, J.D., Sedman, A.J., 1994. An examination of the micromechanics of failure of bone and antler by acoustic emission tests and laser scanning confocal microscopy. Medical Engineering and Physics 16, 203–212. Zioupos, P., Wang, X.T., Currey, J.D., 1996. Experimental and theoretical quantification of the development of damage in fatigue tests of bone and antler. Journal of Biomechanics 29, 989–1002.
Journal of Biomechanics 36 (2003) 1683–1689
Tensile damage and its effects on cortical bone S.P. Kothaa, N. Guzelsub,* b
a Biomechanics and Biomaterials Laboratory, University of Missouri—Kansas City, 650 E. 25th Street, Kansas City, MO 64108, USA Biomechanics Laboratory, University of Medicine and Dentistry of New Jersey, SOM, Tr. #4, 675 Hoes Lane, Piscataway, NJ 08854, USA
Accepted 2 April 2003
Abstract Plexiform bovine bone samples are repeatedly loaded in tension along their longitudinal axis. In order to induce damage in the bone tissue, bone samples are loaded past their yield point. Half of the bone samples from the damaged group were stored in saline to allow for viscoelastic recovery while the others were decalcified. Tensile tests were conducted on these samples to characterize the effects of damage on the mechanical behavior of the organic matrix (decalcified samples) as well as on bone tissue (stored in saline). The ultimate strain of the damaged decalcified bone is 29% higher compared to that of non-damaged decalcified (control) bone. The ultimate stresses as well as the elastic moduli are similar in both decalcified groups. This phenomenon is also observed in other collagenous tissue (tendon and ligament). This may suggest that damage in bone is caused by shear failure of the organic matrix; transverse separation of the collagen molecules or microfibrils from each other. In contrast, there is a trend towards lowered ultimate strains in damaged bone, which is soaked in saline, with respect to control bone samples (not damaged). The damaged bone tissue exhibits a bi-linear behavior in contrast to the mechanical behavior of non-damaged bone. The initial elastic modulus (below 55 MPa) and ultimate strength of damaged bone are similar to that in non-damaged bone. r 2003 Elsevier Ltd. All rights reserved. Keywords: Plexiform bone; Decalcified bone; Shear failure; Damage in type I collagen; Organic matrix
1. Introduction Bone microcracks are most likely manifestations of fatigue damage that develop under repetitive loading conditions (Carter and Hayes, 1977a, b; Schaffler et al., 1990, 1995; Wenzel et al., 1996; Zioupos et al., 1996; Pattin et al., 1996; Boyce et al., 1998; Burr et al., 1998; Zioupos and Casinos, 1998; Fazzalari et al., 1998; Donahue et al., 2000; Yeni and Fyhrie, 2002). Numerous microcracks (o10 mm in length) disperse throughout the bone specimen when they are loaded in tension beyond their yield point (Zioupos et al., 1994; Zioupos and Currey, 1994) along with some larger cracks (Burr et al., 1998; O’Brien et al., 2000). There are no studies that illustrate the nanostructural evolution of damage. Different scenarios of damage initiation leading to *Corresponding author. Department of Osteosciences/Biomechanics, Unviersity of Medicine and Dentistry, PCC Suite 102, 40 East Llaurel Road, Stratford, NJ 08084-1504, USA. Tel.: +1-856-566-2731; fax: +1-856-566-2733. E-mail address: [email protected] (N. Guzelsu). 0021-9290/03/$ - see front matter r 2003 Elsevier Ltd. All rights reserved. doi:10.1016/S0021-9290(03)00169-6
cracks in bone tissue have been suggested (Nicolella et al., 1997; Yeni and Fyhrie, 2002). Damage in the form of cracks can start and develop by tensile failure of collagen molecules/fibrils and/or mineral, by shear failure between collagen molecules/fibrils, and/or by debonding of the organic matrix from the mineral platelets (Nicolella et al., 1997; Yeni and Fyhrie, 2002). A better understanding of the nanoscale damage mechanism in bones due to excess repetitive loads may help physicians to devise treatments for stress fractures more effectively in athletes and army personnel (Macera, 1992; Jones et al., 1993; Ross, 1993; Brill and Macera, 1995; Schaffer et al., 1999; Beck et al., 2000; Givon et al., 2000; Popovich et al., 2000; Marx et al., 2001) as well as in the prediction of fracture risk in osteoporotic bones. Studies of the distribution of stresses and strains within the organic and mineral components may illuminate the nanostructural origin of damage in bone. Cortical bone is considered a short fiber reinforced composite where hydroxyapatite platelets reinforce the organic matrix (primarily type I collagen) (Currey, 1964,
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1969a, b; Carter and Spengler, 1978). When bone is subjected to axial tensile stresses, the organic matrix is subjected to both normal and shear stresses (Currey, 1964, 1969a, b; Carter and Spengler, 1978; Wagner and Weiner, 1992; Jager and Fratzl, 2000; Kotha et al., 2000; Kotha and Guzelsu, 2002) that may cause nanoscale damage in the organic matrix of bone. Further delving into the hierarchical organization of the organic matrix indicates that the organic matrix, which is made up of mainly type I collagen, acts as a composite material itself. The collagen fibrils transfer the tensile forces from one fibril to the next through shear stresses that are generated along the overlap region of the fibrils (Birk et al., 1991; Silver, 1987). Deformation of collagen fibers involves molecular stretching and slippage, fibrillar slippage, and, ultimately, defibrillation (Pins et al., 1997). These mechanisms decrease the fibril diameter and increase the toe region during subsequent tensile testing. Studies of the nanoscale damage mechanism in connective tissues made of type I collagen (tendons or ligaments) have been conducted by mechanically loading these tissues and introducing damage in their ultrastructure (Kastelic and Baer, 1980; Provenzano et al., 2002). Damaged tendons (damaged by loading the tendon up to a certain point and releasing the load) have increased strains to failure compared to nondamaged tendons (Kastelic and Baer, 1980). Provenzano and his co-workers have shown that damaged ligaments (damaged by pulling to a threshold value of 5.14% strain and allowed to recover) demonstrate an increased low-strain region deformation compared to intact samples (increased toe region) (Provenzano et al., 2002). Subfailure damage has been shown to alter the mechanical properties of the anterior cruciate ligament and lengthen the toe region of the force–displacement curve, which result in increased joint laxity (Panjabi et al., 1996). These studies support Kastelic and Baer’s observation that nanoscale damage in the collagen structure increases crimps and results in increased toe region deformation (Kastelic and Baer, 1980). We investigate if these phenomena are also observed in the organic matrix of damaged bone tissue. In this study, we introduce a new method to assay the nanoscale damage phenomena in the organic matrix of bone tissue based on the above studies. We induce tensile damage in cortical bone by repetitively loading it past its yield point. We test the hypothesis that damaged bone tissues exhibit a bi-linear stress–strain behavior while non-damaged bone tissues exhibit a linear stress– strain behavior before their yield point, when calcified samples from these two groups are tested in tension. Furthermore, we test the hypothesis that the ultimate strain of the organic matrix of damaged bone tissue is greater than that of the organic matrix of non-damaged bone tissue when decalcified samples from the two groups are tested in tension.
2. Methods Two-year-old fresh cow femur diaphysis, which were obtained from a local slaughterhouse (Moyer Packing, Sauderton, PA) were used to make tensile mechanical test specimens. The plexiform nature of the samples was confirmed through optical microscopy. Mechanical tests were performed on an Instron 1321B universal testing machine and data was recorded at 60 Hz. The range of the load cell for calcified specimens was 2000 N (resolution of 2.0 N) while for testing decalcified specimens the range was 1000 N (resolution of 1.0 N). In these experiments, 40 samples were divided into four groups. Dumbbell shape mechanical test specimens were made from the lateral and medial cortices of femurs according to ASTM standards (ASTM, 1998; DePaula et al., 2002; with a 2.0 mm 5.0 mm cross-section of the gauge length). Two experimental group samples were repeatedly loaded (eight cycles) in tension past their yield point under stroke control (at a strain rate of 0.0005 s 1) using a triangular waveform to induce damage. The yield point was determined using the 0.2% offset strain method. While they were repeatedly loaded, the average maximum stresses and strains in the bone samples were 110 MPa (110713.8 Mpa–Mean7SD) and 0.79% (0.7970.12%), respectively. Experiments were conducted carefully to prevent the samples from being loaded in compression. Two samples broke during this procedure resulting in nine samples in each of the two damaged experimental groups. The strain in the mineralized samples was measured using an extensometer with a gage length of 12.5 mm. The range and resolution of the extensometer were set to 70.5 and 0.0005 mm, respectively. The recorded stress–strain curves confirmed that the specimens were loaded past their yield point (Fig. 1). Nine damaged and 10 non-damaged samples were decalcified in 0.5 M EDTA solutions for 10 days. Solutions were changed every 2 days (pH=8.3 at 4 C). Each specimen was decalcified in 150 ml solution (Catanese et al., 1999). Sample ends were coated with paraffin to expose a 14 mm rectangular region in the middle of the samples. With this technique, the ends of the samples were kept intact. Decalcification was confirmed as only 0.2% by weight of calcium was eluted in 1 N hydrochloric acid solution as measured by absorptiometry (Sarkar and Chauhan, 1967). These samples were equilibrated for 1 day at 37 C in physiological saline (0.145 M NaCl, pH=7.5) before being mechanically tested. The remaining damaged nine samples and non-damaged intact specimens (10 of them) were immersed in 0.145 M saline at pH=7.5 and 37 C for 8 days. Immersion of damaged bones in saline solution for 8 days allowed the bone samples to recover their viscoelastic deformation after repeated loading
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beyond their yield point. Immersion of bones in saline for a long period of time causes a small amount of bone mineral dissolution, leading to a small reduction in the mechanical properties (Gustafson et al., 1996). Dissolution of bone mineral does not cause a bi-linear stress– strain curve type behavior (DePaula, 2000). In our tension experiment, we compared the mechanical behavior of damaged (experimental calcified group—to allow for viscoelastic recovery) and non-damaged (calcified control—similar equilibration time as experimental group) bones after immersion in saline for 8 days. All four groups were tested in tension under stroke control at a strain rate of 0.0005 s 1. The samples were
Fig. 1. Stress–strain behavior of the same sample tested at a strain rate of 0.0005 s 1. The average maximum stresses and strains in the bones while they are repeatedly loaded are 110.4713.8 MPa and 0.7970.12%, respectively. Note the bi-linear stress–strain behavior of damaged bone after immersion in saline for a period of 8 days. The slope of the upper part of the stress–strain curve of damaged bone is significantly lower at the end of the eighth cycle (Po0:005) due to repeated tensile loading. All values are mean7SD.
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kept wet during the testing and they were loaded to failure in tension. Strains in the decalcified samples were obtained from the increase in displacement between the grips for samples with gage lengths of 12.5 mm (range and resolution of the displacement transducer were 75 and 0.005 mm, respectively). For statistical analysis, one-tailed t-tests were performed to obtain significance.
3. Results The damaged calcified bones had a bi-linear stress– strain behavior before the macroscopic yield point (Fig. 1). These specimens showed a slope transition at about 50–60 MPa. Bi-linear stress–strain behavior is not clearly observable for the 8th cycle of initial loading (Fig. 1, Table 1). Moreover, there was a trend for the damaged bones to exhibit lower ultimate strains (by B30%) in comparison with non-damaged specimens (Fig. 2, Table 1—Columns 5 and 6, Po0:055; post hoc analysis indicated power=0.53 at alpha=0.05). There was a trend towards lower toughness in damaged specimens compared to non-damaged specimens because of lower ultimate strains in damaged specimens and similar ultimate stresses of both groups (Po0:07; post hoc analysis indicated power=0.59 at alpha=0.05). The initial elastic modulus of the control samples and the damaged specimens were similar (Table 1—Column 3). On the other hand, above the 55 MPa stress level, the elastic modulus of the damaged samples was 40% lower than the control samples (Table 1— Column 4, Po0:05; post hoc analysis indicated power=0.996 for alpha=0.05). The damaged specimens that were decalcified showed 29% higher ultimate strains compared with nondamaged decalcified specimens (Po0:05; post hoc analysis indicated power=1.000 for alpha=0.05) (Fig. 3). There were no significant differences in the
Table 1 Elastic moduli, ultimate stress and ultimate strain for the calcified bone samples Groups
Treatments and applied damaged load cycles
Initial elastic modulus below 55 MPa stress level in GPa
Elastic modulus above the 55 MPs stress level in GPa
Ultimate stress in MPa
Ultimate strain (%)
Calcified control samples
Soaked in saline for 8 days
19.673.6
17.172.9a
103.3720.9
2.5770.90c
Calcified damaged samples
First cycle
21.771.6
19.072.5b
Eighth cycle Soaked in saline for 8 days after the eighth cycle
18.372.2 20.273.1
12.472.7b 10.472.4a
106.2716.7
1.7971.01c
Similar alphabets indicate that the values are statistically different from one another (Po0:05).
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(control) versus 0.3270.12 GPa (damaged)} of the 2 decalcified groups (Fig. 3). This indicates that the increased strain in the organic matrix of damageddecalcified bones compared to control-decalcified bones occurs primarily in the toe region.
4. Discussion
Fig. 2. The stress–strain curves for control and damaged bone after being immersed in saline for 8 days. Note the bi-linear stress–strain curve for damaged bone that is not present in control bone. The slopes of the lower part of the two groups are statistically similar (both groups were stored in saline for 8 days). The slope of the upper part of the stress–strain curve of damaged bone is significantly lower than controls (Po0:001). The ultimate strains of damaged bone are significantly different from non-damaged controls (Po0:055) while the ultimate stresses are statistically similar. All values are mean7SD.
Fig. 3. Representative stress–strain curves for decalcified control and decalcified damaged bone. Note the increased ultimate strains in the decalcified organic of the damaged bone as compared with controls (Po0:005). The ultimate strengths of the damaged group are statistically similar to controls. The slopes of the stress–strain curves for both groups are statistically similar in the linear portions. It is noted that three of the samples (two in controls and one in the damaged group) had slight decrements in stresses past the maximum stress. All values are mean7SD.
ultimate strengths {19.3274.25 MPa (control) versus 21.4073.89 MPa (damaged)} or the elastic modulus between 30% and 60% level {0.3270.11 GPa (control) versus 0.2970.13 GPa (damaged)}, and the elastic modulus between 60% and 90% level {0.3670.10 GPa
The presence of damage in bone tissue has been shown to increase the ultimate strain of the organic matrix of the damaged bone as compared to the organic matrix of control bone (Fig. 3). The increase in ultimate strains of the organic matrix in damaged bone was primarily due to changes in the toe region of the stress– strain curve. In collagenous tissue (tendon), the increase in ultimate strains is attributed to a straightening of the crimps as observed with a polarizing microscope (Kastelic and Baer, 1980). We were not able to observe the increased crimps through a polarizing microscope in our decalcified samples. This may be because the organic matrix in the bone tissue is highly disoriented in comparison with the organic structure in tendon. The organic matrix of bone tissue acts to transmit forces, transfer forces among the bone mineral platelets, and prevent bone tissue from failing prematurely as a brittle material. Under excess load, deformation of collagen fibers involves stretching, slippage of laterally adjoining elements and separation (in collagen molecule and/or in collagen fibril levels) and ultimately defibrillation of the fibrils from the overlap regions under shear force transmission (Pins et al., 1997; Sasaki and Odajima, 1996; Sasaki et al., 1999; Christiasen et al., 2000). Provenzano and his co-workers (Provenzano et al., 2002) showed that the stress–strain curve of rat medial collateral ligaments (MCL) that were allowed to recover their viscoelasticity after being loaded above the damage threshold of 5.14%, exhibit an elongated toe region as well as decreased tangential modulus and ultimate stress. Strains of up to 5.14% only increased the elongation in the toe region of MCLs without affecting the elastic modulus or ultimate strengths. Kastelic and Baer also observed a similar behavior in their experiments on rat tail tendons (Kastelic and Baer, 1980). These findings support our findings, wherein, load-induced slippage or separation of some of the laterally joined microfibril structures could reduce the effective diameter of the fibrils and increase the toe region deformation of the collagen structure. In their studies (Provenzano et al., 2002), above 5.14% strains, preloading conditions might be causing lateral defibrillation due to shear forces and also the breakage in the linearly attached fibrils and/or breakage of the fibrils due to axial tension forces, which can alter the elastic modulus and ultimate stress. This behavior seems to be a characteristic of type I collagen as experiments on self
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assembled type I collagen fibrils indicate that a decrease in the diameter of the collagen microfibrils (thinner microfibrils) increases the toe region (low-strain elastic modulus) while an increase in the length of the collagen microfibrils increases the elastic modulus at higher strains as well as the ultimate strength (Christiasen et al., 2000). In a separate study (Yahia et al., 1990), rabbit MCLs subjected to different predetermined strain levels, showed that strains of 10% and above result in numerous broken thin collagen fibers and thick collagen fiber bundles under scanning electron microscopy. Below 10% strain, they did not observe disruption of the collageneous entities. This also indicates that up to certain threshold strain level, lateral separation (defibrillation) could take place in the collagen structure without breakage in linearly attached fibers (breakage across the fibrillar structure) (Christiasen et al., 2000; Provenzano et al., 2002). All these observations suggest that increased crimps are present in the decalcified organic of damaged bones due to decreased lateral association between collagen molecules/microfibrils resulting from a shear failure between collagen molecules/microfibrils as is observed in other collageneous tissues (tendon, ligaments) (Kastelic and Baer, 1980; Provenzano et al., 2002; Christiasen et al., 2000). The dissociation between adjacent microfibrils due to shear damage allows the damaged organic bone matrix to be more crimpled as in damaged tendon (Kastelic and Baer, 1980). This results in increased toe region strains during subsequent mechanical testing. One of the limitations of this study is that the amounts of denatured collagen or the cross-link content were not measured. Our measurements of the mechanical properties of non-damaged decalcified bovine bone are lower than reported by Bowman et al. (1996) for decalcified bovine bone but are within the values reported by Catanese et al. (1999) and Wang et al. (2002) for decalcified human bone tissue. The reason for this discrepancy with the results of Bowman et al. (1996) may be due to their use of preload cycles (up to 5% strain) or the age (not specified) and the location (humerus) of samples. Another limitation is that the number and size of microcracks were not measured in the non-decalcified bones. According to Reilly and Currey (1999), microcracks (o10 mm) were initiated at 0.4% tensile strain in equine bone and there was a considerable growth in microcrack density when tensile strain reached 0.8%. As the strains reached 0.78% in our loading protocol, it was believed that a large number of microcracks (o10 mm) would be present in the bovine bone used in these experiments. The stiffness of bone decreased by 35% (19.4–12.4 GPa) during loading (from first cycle to the eighth cycle) (Fig. 1) and Burr and his co-workers (Burr et al., 1998) have shown larger microcracks after a 15%
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stiffness degradation during bending of a canine femur, a smaller number of larger microcracks may also be present in our damaged samples. The bi-linear behavior of the initial part of the stress– strain curve after the first cycle (different slopes below and above 55 MPa; Fig. 1) are also observed by other researchers (Fondrk et al., 1988; Bonfield and Li, 1967). Fondrk and his co-workers suggested that that the bilinear stress–strain behavior was due to plastic flow and/ or slip behavior or incomplete crack closure (Fondrk et al., 1999). This point of view was supported by the increase in residual volume of bovine and human cortical bone loaded at high-strain amplitudes (Fondrk et al., 1999) and the presence of inelastic strains (Fondrk et al., 1988; Bonfield and Li, 1967). The nanostructural mechanisms responsible for the plastic flow, slip behavior or incomplete crack closure may be related to the damage mechanism within the collagen microfibril such that the presence of mineral within the sheared collagen microfibrils may hold it in an extended configuration, preventing it from reorganization. The mechanical behavior of non-decalcified damaged bone in this study is similar to that described in the literature for bone repeatedly loaded in tension (Zioupos et al., 1994; Jepsen et al., 2001). This behavior is better delineated after the viscoelastic strain is recovered after immersion in saline (8 days in our experiments) (Jepsen et al., 2001). Essentially, the initial elastic modulus of damaged bone is the same as that of non-damaged bone and stress–strain curve becomes curvilinear at higher stresses (>55 MPa in our case). Our results contrast with the results obtained by Carter and Hayes (1977b) in which cortical bone subjected to cyclic bending had lowered residual strength and initial elastic modulus, but agree with studies by Martin et al. (1997) where the residual strength and the initial elastic modulus did not decrease. This may be due to differences in the extent of damage induced in the bone tissue. The technique introduced in this study (decalcifying bone samples after inducing damage), allows us to investigate what happens in the organic bone matrix due to damage in bone tissue. Although we did not directly measure the micro- or macrocracks in our samples, our experimental protocol was similar to that used by other researchers to generate damage in cortical bone. Bone as a composite material is affected by the mechanical properties of both phases. The changes taking place in the organic matrix, which can be investigated with this technique, could be one of the reasons for microfracture initiation in bone samples. As suggested by Nicolella and his co-workers and, Yeni and Fyhrie; shear failure between the collagen molecules/ fibrils, supported by this study, may be one of the main factors starting microcracks in the bone tissue due to excess loading (Nicolella et al., 1997; Yeni and Fyhrie, 2002).
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