Ann Thorac Surg 2002;73:455-459
© 2002 The Society of Thoracic Surgeons
Original article: cardiovascular
Prosthetic replacement of the aorta is a risk factor for aortic root aneurysm development
Natascha Simon-Kupilik, MDa,
Heinz Schima, PhDb,
Leopold Huber, MEb,
Reinhard Moidl, MDa,
Gerhard Wipplingerb,
Udo Losert, MDc,
Ernst Wolner, MDa,
Paul Simon, MD*a
a Department of Cardiac and Thoracic Surgery and LBI for Cardiosurgical Research, University of Vienna, Vienna, Austria
b Institute for Biomedical Engineering, University of Vienna, Vienna, Austria
c Center for Biomedical Research, University of Vienna, Vienna, Austria
Accepted for publication October 17, 2001.
* Address reprint requests to Dr Simon, Department of Cardiac and Thoracic Surgery, University of Vienna, Währinger Gürtel 18-20, A-1090 Vienna, Austria
e-mail: paul.simon{at}univie.ac.at
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Abstract
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Background. Noncompliant prostheses are used in aortic replacement. We hypothesized that this leads to increased distension and wall stress in the aortic root because of the loss of ventriculo arterial coupling.
Methods. Pressure relations in the aortic root caused by changes of aortic elasticity simulating prosthetic aortic replacement were tested in a computer model. We then developed an in vitro model using porcine aortas and performed in vivo validation.
Results. Findings in vitro and in vivo confirmed the predicted changes of the computer model. Pressure amplitude increased significantly by 17% after prosthetic replacement (p < 0.01). Pressure-time differential (Dp/dt) and dicrotic notch pressure amplitude both increased significantly. Echocardiography demonstrated systolic aortic root distension with percentage area change increasing in vitro from 28.2% ± 9.7% to 35.9% ± 10% (p < 0.05) and in vivo from 13.3% ± 3.1% to 24.3% ± 3.1% (p < 0.0001). Aortic root wall stress increased markedly.
Conclusions. Replacement of the aorta with vascular prostheses causes important negative alterations of hemodynamics and increases in wall stress.
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Introduction
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The surgical results of aortic aneurysm repair have greatly improved. Despite continued refinements in surgical technique and postoperative management, patients are at higher risk for late cardiac and aortic morbidity and mortality [1]. Late development of aneurysms in other segments of the aorta has been observed in several patient series [2, 3]. Patient-related factors and surgical technical factors are considered important causes. In a series of patients who had prosthetic replacement of the ascending aorta we found a significant number with aortic root aneurysm development requiring reoperation [4]. We hypothesized that replacement of the ascending aorta with the currently available noncompliant vascular prostheses caused a loss of ventricular-arterial coupling reflected by an increase in aortic root distension, pressure amplitude, and aortic root wall stress, which could contribute to late aneurysm formation. We tested this hypothesis in a computer simulation, in an in vitro experiment, and finally validated the observations in an in vivo experiment.
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Material and methods
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Computer simulations
We first tested our hypothesis using an existing computer model of the complete circulation based on concentrated parameters. This model had been initially designed for interactions between the heart, aorta, and cardiac assist devices, with particular focus on modeling of cardiac and arterial impedances [5]. For this study, segments with standard fluid inertia and resistance, as well as variable elastic and viscoelastic properties for different locations were modeled in the aorta.
The ventricle and aortic valve were followed by a first aortic segment with natural compliance representing the sinus of Valsalva. Four consecutive segments modeled the subsequent prosthesis with compliance elements reduced to one tenth of the natural value. Thereafter, seven segments with natural compliance were arranged, followed by a simplified peripheral impedance (Fig 1). The model is written in the simulation language AGO (Schulz GesmbH, Dortmund, Germany) and runs on a PC586 (Gericom, Linz, Austria).
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Fig 1. Computer model of the aortic root and the adjacent structures. The diagram depicts the structure of the numerical simulator. Each vascular segment is represented by its fluid inertia (I), fluid resistance (R), tissue elasticity (E), and tissue viscoelasticity (Ev, Rv). The left ventricle is modeled by a varying pressure source (PLV), varying compliance (CLV), and myocardial resistance (RLV). The valves are represented by rectifiers (DMITR, DAOR) and leakage resistors (RL); the inflow is provided by left arterial pressure (LAP). The peripheral circulation is simulated by similar parameters IP, RP, EVP, RVP, and a concentrated peripheral resistance RP. For the simulation of the prosthetic graft, the compliances (EG, EVG) are reduced to one tenth of the natural value. (g = graft; r = root; a = aorta; p = periphery.)
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In vitro experimental setup
A modular hydraulic model was used to mimic aortic flow and pressure conditions [6] (Fig 2).
Briefly, the circulation is driven by a piston pump (Vivitro, Victoria, British Columbia, Canada), which provides a computer-controlled flow pattern. A rubber tube, a damping fiber element, and an adjustable resistor mimicked peripheral arterial impedance. We incorporated into this model a fresh porcine aorta that was obtained from a local slaughterhouse. The specimen included the heart and was prepared surgically in our laboratory to contain the intact aortic root with the native aortic valve. The aorta was suspended and held in three positions (proximal arch, descending thoracic aorta beyond the arch, and distal segment of the aorta). It was then connected to the driving chamber using the native valve as the pump outflow valve and was connected distally to the tubing representing the vascular bed. Instrumentation consisted of two Millar micro-tip catheters for hemodynamic measurements with one placed proximal in the aortic root and one just distal to the left subclavian artery. An electromagnetic flow probe (Hellige, Freiburg, Germany) was placed on the distal aorta. In addition a 5-MHz echocardiographic probe was set up in a fixed position to obtain a cross-sectional view of the aortic root at the sinotubular junction, where leaflet motion does not interfere with automatic echocardiographic contour detection (Hewlett-Packard Sonos 2500; Hewlett-Packard, Palo Alto, CA) for measurements of aortic root dimensions. Experiments were performed with eight different aortic specimens. The perfusate used was oil in water suspension. Measurements were performed at a constant flow rate of 3.1 L/minute, and pulsatility was adjusted at baseline to obtain aortic waveforms that resembled physiologic aortic pressure waves. After we obtained baseline measurements, the ascending aorta was wrapped with a commercially available vascular prosthesis (Sulzer Vascutek Ltd, Renfrewshire, Scotland) in the diastolic dimension of the aorta (mean diastolic cross-sectional area 3.8 ± 0.8 cm2) to simulate ascending aortic replacement. Wrapping started beyond the sinotubular junction and continued to the brachiocephalic trunk, with the prosthesis length of 2 cm kept equal in all experiments. Great care was taken to avoid constriction of the aorta. Measurements were then repeated, and each aortic specimen served as its own control.
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Fig 2. Schematic of in vitro model. The thoracic aorta (B) is suspended with the native aortic valve functioning as the outlet valve. Distally the aorta is connected to rubber tubing (C) with adjustable resistance elements (D). The circulation is driven by a motor piston pump (A). E is a reservoir and F represents the mechanical mitral inlet valve.
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In vivo experimental setup
Experimental animals were kept in accordance with published guidelines of care and use of laboratory animals. Validation experiments were performed in five pigs with weights ranging from 18 to 25 kg. After general anesthesia was administered, a left-sided thoracotomy was performed and the pericardium opened and suspended. The aorta was dissected free from the root to the descending thoracic aorta. Instrumentation was the same as in the in vitro model, with two micro-tip catheters inserted in the aorta. One was positioned at the aortic root, the other just distal to the left subclavian artery in the descending aorta. A flow probe was positioned on the trunk of the pulmonary artery. Echocardiographic measurements of the sinus of Valsalva were again made with a 5-MHz probe that was oriented to obtain a cross-sectional view of the aortic root at the sinotubular junction.
After we obtained baseline hemodynamic and echocardiographic measurements, the ascending aorta was wrapped with a 2-cm vascular prosthesis (Vascutec) using the echocardiographically determined diastolic diameter of the aorta as a guideline. Subsequently, the protocol of measurements described above was repeated with the aorta wrapped.
Hemodynamic and echocardiographic measurements
We measured the same variables in the in vitro and in vivo experiments. Proximal, ie, aortic root, pressure and distal aortic pressure and blood flow were recorded continuously and stored digitally for later analysis. Hemodynamic parameters measured included heart rate; systolic, diastolic, and mean pressure; positive and negative pressure-time differential (+dp/dt and -dp/dt); and the maximum and minimum dicrotic notch pressure. The aortic pressure amplitude and the difference of maximum and minimum dicrotic notch pressure were calculated (Fig. 3).
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Fig 3. Automated edge detection. Shown is an example of echocardiographic continuous automated border detection of the aortic root at baseline (a) and after prosthetic replacement of the ascending aorta (b). Changes in cross-sectional area throughout the cardiac cycle are depicted as curve 1. Note the obvious increase in peak systolic area indicating increased distension of the aortic root after replacement of the ascending aorta. Curve 2 is the aortic root pressure again with a marked elevation of pressure amplitude after wrapping.
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Echocardiography of the aortic root was performed with a Hewlett-Packard Sonos 2500 system and a 5-MHz probe. Two-dimensional echocardiography was used to obtain diastolic and systolic diameters and aortic wall thickness for calculation of the following simplified stress index: PR/W = (P x LVID/2)/W (mm Hg), where P represents the blood pressure, R the aortic radius, W the wall thickness, and LVID = left ventricular inner diameter [7]. Acoustic quantification mode [8] was used for automated edge detection (Fig. 3). In this mode a continuous recording of the cross-sectional aortic root area can be performed and is represented as a curve. Percentage fractional area change was calculated from diastolic and systolic areas.
Statistical analysis
At each time point (baseline and after wrapping of the vascular prosthesis), five measurements of each hemodynamic and echocardiographic parameter were made. Two observers who were blinded to each other performed hemodynamic and echocardiographic measurements. The mean of these five measurements was used for further statistical analysis. All data are expressed as mean ± standard deviation. A paired Student’s t test was used to compare baseline measurements with those after aortic wrapping. Statistical significance of differences was assumed at p less than 0.05.
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Results
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The computer model strongly supported our hypothesis of pressure changes after the ascending aorta is stiffened. At a given setting (60 beats per minute, mean aortic pressure 95 mm Hg) the pulse amplitude increased from 44 to 53 mm Hg. After closure of the aortic valve, there were pressure oscillations with an amplitude up to 10 mm Hg and a frequency of 15 Hz. These oscillations were confirmed by careful checking of the stability conditions of the model. They can be explained by the interaction of the fluid inertia within the prosthesis and the elasticity distal and proximal to the prosthesis.
Important hemodynamic changes were also observed in both in vitro and in vivo experiments after wrapping the ascending aorta. In vitro systolic proximal aortic pressure and pressure amplitude increased significantly after prosthetic wrapping, from 103 ± 13 mm Hg to 112 ± 17 mm Hg (p = 0.006) and 58 ± 14 mm Hg to 70 ± 20 mm Hg (p = 0.005), respectively. The amplitude of the maximum and minimum dicrotic pressure increased from 4.8 ± 2.0 mm Hg to 9.3 ± 4.5 mm Hg (p = 0.009). There was a marked increase in maximum pressure-time differential from 281 ± 45 mm Hg/second to 350 ± 77 mm Hg/second (p = 0.006).
The in vivo heart rate and mean arterial pressure remained constant. The changes in the in vitro model were less pronounced but could still be detected. Systolic aortic pressure increased after aortic wrapping from 127 ± 11 mm Hg to 134 ± 15 mm Hg, and the pressure amplitude increased from 37 ± 9 mm Hg to 46 ± 15 mm Hg. This is a 17% increase in pressure amplitude after prosthetic wrapping. There was a significant increase in maximum pressure-time differential from 898 ± 373 mm Hg/second to 1353 ± 350 mm Hg/second (p = 0.03), a change of 36%. In addition the amplitude of the maximum and minimum dicrotic pressure increased significantly from 4.4 ± 1.4 mm Hg to 9.4 ± 2.8 mm Hg (p = 0.002) (Fig 4).
This 53% increase reflects pressure oscillations within the aortic root.
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Fig 4. Example of pressure patterns at the aortic root in vivo. An increase in the pressure amplitude and a significant increase of the amplitude of the maximum and minimum dicrotic notch pressure without (A) and after (B) graft reinforcement are shown. (AoP = aortic pressure).
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Echocardiographic data showed, both in vitro and in vivo, the effects of the hemodynamic changes on the aortic root after prosthetic wrapping (Table 1).
Diastolic aortic root cross-sectional area remained unchanged both in vitro and in vivo. However, systolic area increased after wrapping, with an increase in percentage fractional area change reflecting increased systolic root distension (Fig 3). The wall stress index increased significantly in vitro and in vivo by 22% and 16%, respectively.
The in vitro and in vivo observed hemodynamic changes confirmed the predictions of the computer model, which also revealed significant pressure oscillations after valve closure with frequencies higher than 15 Hz at amplitudes of up to 10 mm Hg. These oscillations were confirmed by careful check of stability conditions and result from the inertia (inductivity) in the prosthesis without the natural buffering by elasticity.
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Comment
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Our in vitro and in vivo data demonstrate that replacement of an aortic segment with the currently available stiff, noncompliant vascular prostheses leads to significant hemodynamic changes in the aortic root. This was shown by a significant increase in systolic cross-sectional area of the aortic root and a significant increase in calculated wall stress index. This supports the work of Schulz and colleagues [9] who found in their computer model that the aortic input impedance significantly increased after prosthetic replacement of the aorta. Similarly, aortic input impedance was shown to be dramatically increased by 255% in an experimental model of extra-anatomic aortic bypass [10].
Several limitations of the study must be considered in the interpretation of our data. First, the in vitro setup reflects artificial conditions, even if one tries to obtain aortic pressure wave forms that are similar to those obtained under physiologic conditions, because of the different impedance of the pulsatile heart pump and the simplification of the impedance of peripheral arteries. Nevertheless, the same effects were documented in the in vivo setting, albeit slightly less pronounced (ie, with higher damping). The intact organism may be able to compensate for some of the changes seen in vitro. Second, the porcine aorta used was highly compliant, whereas the human aorta loses compliance with age and in certain disease states. Therefore the changes seen in our experiment may be less marked in patients with aneurysms, although this has yet to be determined.
Nevertheless, in several series it has been shown that patients who had aortic procedures have significant cardiac and aortic long-term morbidity and mortality [2, 11, 12]. Schepens and associates [13] demonstrated clearly that there is a significantly lower survival rate than would be expected when compared with a background population. The increase in input impedance leads to ventricular hypertrophy and might contribute to the observed increase in cardiac mortality rate [14, 15]. Aortic rupture is a well-documented late complication, and life-long radiographic surveillance of these patients is recommended. We observed an increased incidence of sinus of Valsalva aneurysms after replacement of the ascending aorta [4], which was particularly prevalent in patients who had aortic dissection as their underlying pathology but was also seen in patients with nondissecting aneurysms. The increase in systolic distention of the aortic root and in wall stress might precipitate the development of late aneurysms, particularly in the presence of residual aortic wall pathology and in patients with poorly controlled blood pressure. Moreover, the enhanced kinetic energy of blood due to the high frequency contents of pressure and pressure oscillations could increase the mechanical load on the aortic valve. Ergin and associates [16] found that aggressive treatment of aortic root pathology in patients with aortic dissection with composite graft replacement improved their event-free long-term survival significantly. In many patients the aortic valve itself is morphologically normal, and surgical techniques of aortic root replacement sparing the native valve have been described [17, 18] with excellent results [19, 20].
We conclude from our data that, because of the lack of elasticity of currently available prosthetic materials, prosthetic replacement of the ascending aorta may in itself be a risk factor for late aneurysm formation in the sinus of Valsalva. Patients with poorly controlled blood pressure and residual aortic pathology can be expected to be at increased risk. Further improvement in the prognosis of these patients requires aggressive surgical treatment of all pathology in the aortic root at the primary procedure and strict blood pressure control during follow up. We will use the data obtained here to develop prosthetic materials with more natural compliance.
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Abstract
Introduction
