- Open Access
Characterization of mechanical behavior of a porcine pulmonary artery strip using a randomized uniaxial stretch and stretch-rate protocol
© Jhun and Criscione; licensee BioMed Central Ltd. 2008
- Received: 24 April 2007
- Accepted: 23 January 2008
- Published: 23 January 2008
Much of the experimental work in soft tissue mechanics has been focused on fitting approximate relations for specific tissue types from aggregate data on multiple samples of the tissue. Such relations are needed for modeling applications and have reasonable predictability – especially given the natural variance in specimens. There is, however, much theoretical and experimental work to be done in determining constitutive behaviors for particular specimens and tissues. In so doing, it may be possible to exploit the natural variation in tissue ultrastructure – so to relate ultrastructure composition to tissue behavior. Thus, this study focuses on an experimental method for determining constitutive behaviors and illustrates the method with analysis of a porcine pulmonary artery strip. The method characterizes the elastic part of the response (implicitly in terms of stretch) and the inelastic part in terms of short term stretch history (i.e., stretch-rate) H t 2, longer term stretch history H t 1, and time since the start of testing T.
A uniaxial testing protocol with a random stretch and random stretch-rate was developed. The average stress at a particular stretch was chosen as the hyperelastic stress response, and deviation from the mean at this particular stretch is chosen as the inelastic deviation. Multivariable Linear Regression Analysis (MLRA) was utilized to verify if H t 2, H t 1, and T are important factors for characterizing the inelastic deviation. For acquiring H t 2and H t 1, an integral function type of stretch history was employed with time constants chosen from the relaxation spectrum of an identical size strip from the same tissue with the same orientation. Finally, statistical models that characterize the inelasticity were developed at various, nominal values of stretch, and their predictive capability was examined.
Inelastic deviation from hyperelasticity was high (31%) for low stretch and declined significantly with increasing stretch to a nadir of 3.6% for a stretch of 1.7. The inelastic deviation then increased with increasing stretch at the same point in the stress-strain curve where stiffness began to increase strikingly. MLRA showed that T is a major inelastic parameter at low deformation. For moderate and high deformations, H t 2and H t 1were dominant.
A randomized uniaxial testing protocol was applied to a strip of porcine pulmonary artery to characterize the elasticity and inelasticity of a soft tissue. We were successful in determining the elastic response and the factors that gave rise to the inelastic deviation. This investigation seeks methods to better define, phenomenologically, the elastic and inelastic behavior of soft tissues.
- Constitutive Behavior
- Stretch Ratio
- Relaxation Spectrum
- Inelastic Behavior
- Partial Regression Coefficient
A great challenge in biomechanics, and its related disciplines such as mechanobiology, is the determination of mechanical constitutive behaviors for tissues that undergo high strain. Even for homogeneous, single phase, rubber-like materials with behavior that is nearly isotropic and hyperelastic, the determination of constitutive behaviors has been ambiguous and problematic . For multi-phase, heterogeneous tissues with behavior that is anisotropic and visco-elastic, much theoretical and experimental work remains , and this work is a small step toward improving experimental methods to delineate tissue behavior.
There are, of course, approximate constitutive relations for most tissue types that have been validated with fits to experimental data and that have enabled modeling applications in biomechanics. Rather than integrate tests from multiple specimens to yield a generalized relation, this work is focused on a methodology to study individual specimen variations. Our reason for doing so is to exploit the natural variation in tissue structure to better understand the structure-function relationship. Each specimen has differing amounts of elastin, collagen, water, etc.; and with experimental methods that delineate rather than aggregate behavior, we seek to relate ultrastructure parameters to tissue response parameters. The ultimate goal of this work, albeit decades away, is to determine how ultrastructure maps to tissue behavior.
Toward this end, we developed a random stretch and random stretch-rate protocol for a one-dimensional sample (longitudinal strip from a porcine pulmonary artery). The average stress at a particular stretch is defined as the elastic or hyperelastic part of the response and the deviation from this average is defined as the inelastic part of the response. The inelastic response is assumed to be a function of three variables: stretch rate related, short tem stretch history H t 2; longer term stretch history H t 1; and time T since the start of the test protocol. At a particular stretch value, the inelastic response is assumed to be a linear function of these variables and dependence is found by Multivariable Linear Regression Analysis (MLRA). A randomized testing protocol eliminates any potential correlation among H t 2, H t 1, and T that conventional cyclic testing cannot. This randomness enables the use of MLRA.
Hoffman and Grigg have introduced the advantage of application of random stress stimuli to the soft tissue . The difference here is that stretch herein, rather than stress therein, is the independent variable that is randomized. Moreover, we seek to define and determine the elastic and inelastic parts of the response at particular chosen values of strain rather than fit behavior for the range tested. Though, there are many time-dependent theories that can capture or fit inelastic behaviors, such as, linear/nonlinear viscoelastic [4–10], poroviscoelastic [11, 12], and biphasic [13–16]., this study is focused on experimental methodology for hyperelastic constitutive behaviors.
Remark on Hyperelasticity
We often mix the terms elastic and hyperelastic when referring to material behavior and certain aspects of the stress response. Conceptually, one can think about elasticity (i.e., stress as a function of stretch or strain) as different than hyperelasticity (i.e., strain energy as a function of stretch or strain); yet in fact, elastic behavior must be hyperelastic – or else it violates the first law of thermodynamics. This is easily shown by considering a closed, cyclic process involving an elastic material. An elastic material does not have viscous losses, and the work of the reverse path (or cycle) is the opposite of the forward path (or cycle). This is a direct result of the stress-power law. If the work done is zero for all closed paths, then there exists a state function (strain energy) that represents the storage, in a conservative manner, of work done on the material. This is hyperelasticity wherewith the constitutive law can be defined by its strain energy function and the stress is given by the derivative of the strain energy with respect to the strain. If the work done is not zero for all closed paths (i.e., non-hyperelastic), then choose one such path and go around the cycle in the direction that produces work, and then build a perpetual motion machine of the first kind. It is quite easy (in two and three dimensions) to fit an elastic model that violates the first law ; and yet it is impossible for a hyperelastic model to violate the first law. To be explicit: 1) if an elastic model satisfies the first law then it can be expressed in a hyperelastic form, and 2) if an elastic model cannot be expressed in a hyperelastic form, then it violates the first law. When we refer to elastic behavior as hyperelastic, we are only excluding behavior that violates the first law of thermodynamics.
This stretching protocol was accomplished with custom LabVIEW (National Instruments, Inc) codes. The stretch values were chosen to cover the linear and non-linear portions of the stress-stretch response of the tissue, without damaging the tissue. The stretch-rate range was determined by the upper and lower bound for the speed of the actuators (CMA 25CCCL DC servo motor, max speed = 0.4 mm/s, Newport, Inc). The actuators were controlled by the Universal Motion Controller/Driver (ESP7000, Newport, Inc). The total testing time was 5600 seconds.
The pulmonary artery from an adult swine was obtained from a local slaughter house immediately after slaughter. Two identical strips (25.4 mm × 2.5 mm × 2 mm) were cut from the porcine pulmonary artery (one was used for the stretch protocol and one for the relaxation analysis), preserved in chilled phosphate buffered saline (PBS) and tested at room temperature (approximately 23°C). We mounted a specimen by using custom-built alligator type grips and no slip was detected during the test. The stretch ratios were calculated based on a gauge length which is a grip-to-grip displacement. (In general, grip-to-grip strain is greater than average strain measured by optical strain measurement because of edge effects and local deformation at the tip of the grip. Since the length of our specimen is relatively long relative to its width, 10:1; and since this work is focused on archiving and discussing the methodology as opposed to archiving the behavior of a particular, now decomposed, sample, we neglected edge effects.) The average force at a particular stretch was defined as the hyperelastic response because it would be the best guess for the response if the stretch rate and stretch history were unknown or neglected – as per the assumption of elasticity. Inelastic deviation was defined and measured as the deviation from the average force at each nominal stretch ratio.
To quantify inelastic behavior, we assumed the functional form shown in equation (2) where F M and F W are measured and average force, respectively at a particular stretch. F M = F W ± D(H t 2, H t 1, T)
To explicitly represent the limits of experimental methodologies, we included an unknown factor D e that represents additional inelastic deviation (or random error) beyond that which is represented by D Ht 2, D Ht 1, and D T or deviation due to H t 2, H t 1, and T, respectively. Hence, D = D Ht 2+ D Ht 1+ D T + D e
F(t) = 6.1 + 1.5e-t/400+ 0.5e-t/30+ 1.3e-t/10+ 0.5e-t
Multiple linear regression analysis
The partial regression coefficients , j = 1, 2 and 3 describe how the independent variables T i , (H t 2) i , and (H t 1) i affect the dependent variable D i . We minimized e i , the errors in fitting, by using a least squares method (LSM) and so obtained and . To avoid scale-dependency in the partial regression coefficients, we normalized each data set by their magnitudes. Upon fitting the , we checked , the adjusted coefficient of determination to observe how well the derived sample regression line describes the observed variable D i and then examined the significance of each using t-test. Successive iterations of t-test were performed whenever there was/were insignificant variable(s) for a particular stretch. Since we tested a single specimen, we set the level of significance α as 0.1 which represents the 90% probability that the partial regression coefficient is likely to be contained within the interval. In general, a level of significance α is preferably set as 5%.
Lastly, to assess the goodness of fit and search for further systematic behavior, we calculated the deviation of the predication, D i - , which is the difference between observed deviation of a data point D i and point forecasting value obtained by the statistical model for the particular, nominal stretch value. If the deviation of the predication has systematic variation, then it can reasonably be said that there must be other factors (or perhaps higher orders of H t 2, H t 1, and T) that induce inelastic deviations from hyperelasticity.
Absolute and fractional uncertainties for each stretch ratio.
F W [mN]
Multivariable linear regression models for the tissue (t c is for two-tailed test).
t c(α/2;n-k-1), α = 0.1
This study characterizes the elastic (or mean) response and the inelastic (or deviation from the mean) response of a porcine pulmonary artery strip that underwent high strain with random stretch and stretch-rate prescribed. The inelastic behavior was most prominent for small deformations and primarily dependent on T, the time since the start of testing. The inelasticity at λ = 1.1 was almost 850% greater than that of λ = 1.7, and 400% greater than that of λ = 1.9. This finding can provide one reason for much ambiguity in the hyperelastic modeling for soft tissues at small deformations. The most important configuration for elasticity (i.e., the stress free reference configuration) is in the deformation region wherein inelasticity predominates. How the sample was prepared and any significant stretch history during dissection are likely to affect greatly the reference configuration and any subsequent constitutive relation that depends on the reference configuration. At the highest deformation λ = 1.9, inelastic deviation was increased again, and it is significant that in this specimen, this stretch is at the point where the stiffness starts to increase greatly. We suspect that this break in trend is due to complex microstructural coupling of fibers in radial and circumferential directions and some reorganizing of the fibers with the high deformation. Except for low deformation, the inelastic deviations for the other cases are less than 10%. Thus, an elastic model is mostly, say 90%, accurate for this strip of pulmonary artery.
MLRA clearly indicates that H t 2, H t 1, and T are determinant factors that induce inelastic behavior, but they are not equally influential throughout the stretch domain. Inelasticity of the tissue was dominated by different factors at different stretches, and it was systematic rather than random. T was an influential factor for all deformations, but particularly for low deformations. As the deformation increases, however, stretch rate is the most influential for the inelastic response. Given that our protocol randomized the stretch, it is not very significant that the deviation of the predication appears random or unsystematic after we correct for deviations due to H t 2, H t 1, and T.
In summary, we found that the stretch rate, stretch history, and time are important factors that give rise to the inelastic mechanical response of a strip of pulmonary artery as evident in a randomized uniaxial stretch-controlled stretch test. Their contributions were not the same for all stretches. This study was limited to one soft tissue specimen at room temperature; however, it is the opinion of the authors that one must be able to define the phenomenological behavior of one specimen before it is possible to define relations that represent the aggregate or typical behavior of a tissue. For example, if one could not measure the elastic modulus of a single sample of steel wire, one could not measure the average elastic modulus of steel. Albeit generalized, constitutive relations for particular tissues are mostly accurate and in use already; there is much ambiguity in their definitions. The focus of this work is determining the behavior of one sample, and in so doing, we hope to eventually define quantitatively how specimens are similar, and more importantly for exploiting natural variations in ultrastructure, how they differ.
The American Heart Association (grant 0435210N to JCC), Texas Engineering Experiment Station (TEES) and Department of Biomedical Engineering at Texas A&M University are greatly acknowledged for financial contribution and use of facilities.
- Criscione JC: Rivlin’s representation formula is ill-conceived for the determination of response functions via biaxial testing. Journal of Elasticity 2003, 70: 129–147. 10.1023/B:ELAS.0000005586.01024.95MathSciNetView ArticleGoogle Scholar
- Criscione JC: Kinematics framework optimized for deformation, growth, and remodeling in vascular organs. Biomechanics and modeling in mechanobiology 2007.Google Scholar
- Hoffman AH, Grigg P: Using uniaxial pseudorandom stress stimuli to develop soft tissue constitutive equations. Annals of biomedical engineering 2002,30(1):44–53. 10.1114/1.1432689View ArticleGoogle Scholar
- Braidot AA, Deiber JA: Linear viscoelastic model of a maturing gelatin solution. Biorheology 1999,36(3):267–284.Google Scholar
- Decraemer WF, Maes MA, Vanhuyse VJ, Vanpeperstraete P: A non-linear viscoelastic constitutive equation for soft biological tissues, based upon a structural model. Journal of biomechanics 1980,13(7):559–564. 10.1016/0021-9290(80)90056-1View ArticleGoogle Scholar
- Ehlers W, Markert B: A linear viscoelastic biphasic model for soft tissues based on the Theory of Porous Media. Journal of biomechanical engineering 2001,123(5):418–424. 10.1115/1.1388292View ArticleGoogle Scholar
- Ledoux WR, Meaney DF, Hillstrom HJ: A quasi-linear, viscoelastic, structural model of the plantar soft tissue with frequency-sensitive damping properties. Journal of biomechanical engineering 2004,126(6):831–837. 10.1115/1.1824133Google Scholar
- Rice DA, Yang TY: A nonlinear viscoelastic membrane model applied to the human cervix. Journal of biomechanics 1976,9(4):201–210. 10.1016/0021-9290(76)90005-1View ArticleGoogle Scholar
- Sanjeevi R, Somanathan N, Ramaswamy D: A viscoelastic model for collagen fibres. Journal of biomechanics 1982,15(3):181–183. 10.1016/0021-9290(82)90250-0View ArticleGoogle Scholar
- Shen F, Tay TE, Li JZ, Nigen S, Lee PV, Chan HK: Modified Bilston nonlinear viscoelastic model for finite element head injury studies. Journal of biomechanical engineering 2006,128(5):797–801. 10.1115/1.2264393View ArticleGoogle Scholar
- DiSilvestro MR, Suh JK: A cross-validation of the biphasic poroviscoelastic model of articular cartilage in unconfined compression, indentation, and confined compression. Journal of biomechanics 2001,34(4):519–525. 10.1016/S0021-9290(00)00224-4View ArticleGoogle Scholar
- Haider MA, Schugart RC: A numerical method for the continuous spectrum biphasic poroviscoelastic model of articular cartilage. Journal of biomechanics 2006,39(1):177–183. 10.1016/j.jbiomech.2004.10.037View ArticleGoogle Scholar
- Barocas VH, Tranquillo RT: An anisotropic biphasic theory of tissue-equivalent mechanics: the interplay among cell traction, fibrillar network deformation, fibril alignment, and cell contact guidance. Journal of biomechanical engineering 1997,119(2):137–145. 10.1115/1.2796072View ArticleGoogle Scholar
- Barocas VH, Tranquillo RT: A finite element solution for the anisotropic biphasic theory of tissue-equivalent mechanics: the effect of contact guidance on isometric cell traction measurement. Journal of biomechanical engineering 1997,119(3):261–268. 10.1115/1.2796090View ArticleGoogle Scholar
- Cohen B, Chorney GS, Phillips DP, Dick HM, Mow VC: Compressive stress-relaxation behavior of bovine growth plate may be described by the nonlinear biphasic theory. J Orthop Res 1994,12(6):804–813. 10.1002/jor.1100120608View ArticleGoogle Scholar
- Mow VC, Kuei SC, Lai WM, Armstrong CG: Biphasic creep and stress relaxation of articular cartilage in compression? Theory and experiments. Journal of biomechanical engineering 1980,102(1):73–84.View ArticleGoogle Scholar
- Criscione JC, Sacks MS, Hunter WC: Experimentally tractable, pseudo-elastic constitutive law for biomembranes: I. Theory. Journal of biomechanical engineering 2003,125(1):94–99. 10.1115/1.1530770View ArticleGoogle Scholar
- Reddy JN: An introduction to the finite element method. 2nd edition. New York , McGraw-Hill; 1993:xix, 684 p..Google Scholar
- Taylor JR: An introduction to error analysis : the study of uncertainties in physical measurements. 2nd edition. Sausalito, California. , University Science Books; 1997:xvii, 327 p..Google Scholar
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