DYNAMIC MECHANICAL PROPERTIES OF ARTERIAL AND VENOUS GRAFTS USED IN CORONARY BYPASS SURGERY
Dimosthenis Mavrilas1*, Theodora Tsapikouni*, Dimitrios Mikroulis+, Grigorios Bitzikas+, Vassilios Didilis+, Kosmas Tsakiridis+, Fotis Konstantinou+ and Georgios Bougioukas+ *
Biomedical Engineering Laboratory, Dept. of Mechanical Engineering & Aeronautics, University of Patras, 26500 RION, PATRAS, GREECE. e-mail:
[email protected] + Dept. of Cardiothoracic Surgery, University Hospital of Alexandroupolis, Demokritus University of Thrace, ALEXANDROUPOLIS, GREECE. Key words: mechanical properties, saphenous vein, internal mammary artery, bypass grafts, hysteresis, elastic modulus. Abstract. In this work we studied the frequency dependence of the dynamic mechanical characteristics of saphenous vein (SV) and internal mammary artery (IMA) grafts. Rectangular longitudinal strips from 14 patients were tested under cyclic uniaxial tensile loading in the frequency range of 0.1-20 Hz, at 37oC in wet conditions. The dynamic mechanical parameters (the storage modulus ES and the hysteresis ratio h (loading/loop area)) together with the collagen phase modulus EH were computed as a function of frequency. The results showed that in all graft types ES and EH varied with frequency in the range 0.5–10 Hz, presenting a maximum in the neighboring of 1 Hz. The hysteresis ratio h was increased in the frequency range 1– 20 Hz. It seems from the results that the physical resonance frequency of the components of the tissue responsible for their elastic behavior may lay in the range around 1 Hz, while that for the viscous behavior in the range of 20 Hz or more. Early clinical outcomes of both grafting were studied in parallel. In one - year postoperative period the follow - up (clinical examination, electrocardiography, echocardiography and stress test) did not reveal any sign of graft occlusion or severe stenosis except one perioperative infraction but without any correlation to the graft quality. 1. INTRODUCTION Internal mammary artery (IMA) and saphenous vein (SV) grafting have been widely used for myocardial revascularization. The long-term clinical superiority of IMA over SV grafts has been documented [1,2] with the last one having higher postoperative occlusion rates. In their new anatomic position both grafts bear a different mechanical loading profile than in their physical ones. Coronary flow differs from the flow in internal thoracic arteries and much more from the flow in the veins of lower extremities. Structure and size of coronary arteries are quite different from that of arterial and venous conduits used for bypass grafting [3-5]. These differences have as result the long–term occlusion of grafts caused from thrombosis, intimal hyperplasia and atherosclerosis, with SV grafts to be occluded in much more extend
than IMA grafts [2]. Despite their poor longevity, however, SV grafting remains an alternative when IMA grafting is not available or more coronary vessels demonstrate critical stenosis. The knowledge of flow – pressure interaction with the structure of cardiovascular tissue is still very limited because of the difficulty of performing accurate measurements in vivo, either in humans or in animal models [6]. The physical characteristics of the various vascular tissue materials, as their structure and mechanical performance, are closely related to the mechanical loading imposed, to match the special circulatory demands of their anatomic position. A better understanding of these characteristics could raise the possibility of improving the patency of vascular grafts. The arterial and venous wall consists of three layers: the intima, media and adventitia, From a technological point of view the various vascular graft tissues are structured as multilaminate reinforced composite materials. They exhibit a hardly anisotropic, non–linear, viscoelastic mechanical response to the applied load at strains very high compared to that, which the linear viscoelastic theory is usually applied. Their behavior under mechanical stress is a combination of elastic and viscous components, the exact modeling of which is attempted by numerous approaches [7]. However, to the authors’ opinion, no mathematical or physical generalized models of soft tissues can applied satisfactory for a broad range of either frequency or stressstrain levels. Tissue viscoelastic characteristics, such as storage modulus (ES) and hysteresis are depended on the frequency, the value, the shape and the history of the cyclic mechanical loads applied on the vessel walls. Because of anisotropy the tissue responses differently in radial, longitudinal and circumferential directions. Although the total frequency of the complicated pressure and flow waves lay normally in the range 0.5–2.5 Hz, the composing harmonic frequencies may exceed much further this range, approaching possibly the values of physical resonance of the vessel wall material [8]. So, when new grafts are evaluated prior to use, extended mechanical tests should be performed to investigate their mechanical performance, the specific requirements of which are strongly depended on the certain clinical application of the grafts. In the present work we investigated the dynamic mechanical characteristics of IMA and SV grafts used for bypass surgery under the application of uniaxial cyclic tensile loading of vascular wall tissue strips in the longitudinal direction. We studied and compared the viscoelastic material characteristics, like the storage modulus and hysteresis, as well the tissue high elastic modulus (modulus of the collagen phase) and their dependence on the frequency of the applied load. 2. MATERIALS AND METHODS The tested materials were fresh vascular tissues from IMA and SV, obtained from grafts used for bypass surgery. Fourteen patients, 13 male and 1 female, 47-72 years old were operated on for bypass grafting with 2,9 grafts per patient average. The grafts used were autologous IMA and SV. Small graft segments were cut during the preparation prior to implantation, stored in iced normal saline and sent for mechanical testing. The tests were performed within 24 hr. The cylindrical segments were cut longitudinally, flattened and divided into rectangular longitudinal strips 2.5 x 15-20 mm. The strips were secured carefully by their narrow edges in special designed
Plexiglas grips of a custom made tensile testing machine, leaving a 5–7 mm free length between them. This machine was equipped with an electrodynamic vibrator system (Link–Dynamics V201) which actuated the moving grip. A manually operated screw mechanism permitted the movement of the vibrator for initial permanent elongation of the tissue strips during testing. The moving grip was mechanically engaged with the moving light – weight core of an LVDT elongation transducer (Lucas – Schaevitz DC-EC 250) while its coil was engaged with the body of the vibrator, so that LVDT responded to the vibratory movement of the moving grip and not to the initial elongation. The fixed grip of the testing machine was connected to a 500 N piezoelectric force transducer–charge amplifier system (Kistler 9203-5006). A closed loop circulatory system including a water bath permitted the continuous isothermal moistening of the tissue samples with normal saline at 37oC during testing. A schematic diagram of the testing apparatus is shown in fig. 1. An important step in tensile testing of soft tissues is the determination of their initial dimensions: the “free” length and the cross-sectional area. As vessel wall tissues in physiologic conditions are in a static prestressed state during pressure loading [9] it is important to reproduce this state by preloading the tissue strips prior the final tensile testing. This preloading was performed by the manual movement of the vibrator until a length (referred as initial “free” length (lo) of the tissue strip) at which a slight elongation responded to a signal in force, with both transducers being in high sensitivity. This situation is referred consequently as “zero loading”. Cyclic uniaxial tensile testing was performed in the frequency range of 0.1-20 Hz to minimize possible mechanical noise presented at higher frequencies. Elongation range was varied from the “zero loading” to a value great enough so that the tissue was deep in the characteristic high slope linear region [10] (referred as “collagen modulus region”). Force, elongation and time data were collected via an analog to digital (AD) converter and a PC system with sampling frequency variable, modified at each testing frequency to be great enough for the collection of 100 sets of measurements per loading cycle. The dimensions of the tissue strips (initial length l0, width and thickness) were determined non-contact, by analysis of digital images obtained at zero loading with the use of a high resolution closed circuit video-camera (Sony CCD Iris) and a PC installed video grabber system. Lagrangian stress (σ = F/A0) and strain (ε = ∆l/l0) were computed from load–elongation and dimensions’ data. Experimental stress - strain data were fitted mathematically by Fourier analysis. At the subsequent synthesis, the use of the first three sine cosine terms of the Fourier series were found to satisfactory fit the experimental data. For each tissue strip the computed Fourier mathematical periodic functions were used to obtain stress–strain loading–unloading loop diagrams as that shown in fig. 2, from which the dynamic mechanical parameters such as the storage modulus ES, the collagen phase modulus EH and the hysteresis ratio h were computed. ES was computed from the slope of the straight line from zero to the maximum stress σmax (secant modulus). EH was computed from the slope of the linear part of the loading phase (tangential modulus) at high strains, characteristic of the mechanical contribution of collagen reinforcement of soft tissue [10]. Hysteresis ratio h was computed after numerical integration of the periodic Fourier mathematical functions found, separately for loading and unloading phase as: h = (loading – unloading area)/(loading area) or
h = Loop / loading area. Early clinical outcomes of all patients were studied in parallel. The follow-up included clinical examination and electrocardiogram per three months and echocardiography and stress-test per six months. During one-year follow-up period no sign or finding of graft occlusion or severe stenosis revealed, except one perioperative infraction, but without any correlation to the graft quality. The gold standard for checking the graft patency is the coronary angiography. We did not perform this examination, because none of our patients had any indication for an intervention investigation. 3. RESULTS Figure 3 presents the change of the storage modulus ES with frequency in the range 0.1-20 Hz. At low frequencies (f < 0.5 Hz) the mechanical response looks rather frequency insensitive. In the range of 0.5-1Hz ES is increased (the tissue becomes stiffer) reaching a maximum value and consequently decreased (tissue softening) again up to 5 Hz remaining further at the initial low frequency level at the range 5-20 Hz. Figure 4 presents the change of collagen modulus EH with frequency. At low and high frequency ranges this modulus is in the same level. However at middle range (110 Hz) maximum values are observed (stiffening and softening of the tissue), as in the case of ES, although the maximum EH of IMA and SV are observed at different frequencies (0.5 and 1 Hz respectively). Figure 5 presents the change of hysteresis ratio h with frequency. h looks to be frequency insensitive at 0,1-1 Hz, gradually raised in the range of 1 - 20 Hz until the maximum tested frequency of 20 Hz. It seems that the tissues exhibit a more viscous behavior at frequencies near 20 Hz. In all the three figures, modulus values are mean ± SE. No statistical differences were found between all the three graft types (SVP, SVD, IMA) for each mechanical parameter at every one frequency measured (One way ANOVA, p
