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Ann Thorac Surg 2010;89:1981-1988. doi:10.1016/j.athoracsur.2010.03.002
© 2010 The Society of Thoracic Surgeons
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Original Articles: Adult Cardiac

Comparison of Porcine Pulmonary and Aortic Root Material Properties

Peter B. Matthews, BS, Ali N. Azadani, PhD, Choon-Sik Jhun, PhD, Liang Ge, PhD, T. Sloane Guy, MD, Julius M. Guccione, PhD, Elaine E. Tseng, MD*

Department of Surgery, University of California at San Francisco Medical Center and San Francisco Veterans Affairs Medical Center, San Francisco, California

Accepted for publication March 1, 2010.

* Address correspondence to Dr Tseng, UCSF Medical Center, Division of Cardiothoracic Surgery, 500 Parnassus Ave, Ste W405, Box 0118, San Francisco, CA 94143-0118 (Email: elaine.tseng{at}ucsfmedctr.org).


   Abstract
Top
 Abstract
Introduction
 Material and Methods
 Results
 Comment
 Acknowledgments
 References
 
Background: The pulmonary autograft remodels when subjected to systemic pressure and subsequent dilation can lead to reoperation. Inherent material property differences between pulmonary and aortic roots may influence remodeling but are currently unknown. The objective of this study was to determine stiffness across a wide range of strain and compare nonlinear material properties of corresponding regions of native aortic and pulmonary roots.

Methods: Tissue samples from porcine aortic and pulmonary roots—sinuses and supravalvular artery distal to the sinotubular junction—were subjected to displacement-controlled equibiaxial stretch testing. Stress-strain data recorded were used to derive strain energy functions for each region. Stiffness from low to high strains at 0.15, 0.3, and 0.5 strain were determined for comparisons.

Results: Aortic and pulmonary roots exhibited qualitatively similar material properties; both had greater nonlinearity in the sinus than supravalvular artery. The pulmonary artery was significantly more compliant than the ascending aorta both circumferentially and longitudinally throughout the strain range (p < 0.03), except at high strain circumferentially (p = 0.06). However, no differences in stiffness were seen circumferentially or longitudinally between pulmonary and aortic sinuses (p ≥ 0.3) until high strain, when the pulmonary sinuses were significantly stiffer (p < 0.05) in both directions.

Conclusions: Differences in stiffness between porcine aortic and pulmonary roots are regionally specific, supravalvular artery versus sinus. These regional differences may impact the mode of remodeling to influence late autograft dilation.


   Introduction
Top
 Abstract
 Introduction
Material and Methods
 Results
 Comment
 Acknowledgments
 References
 
The Ross operation replaces a diseased aortic valve with a pulmonary autograft through three implantation methods: subcoronary, root inclusion, or full root replacement, where the full root is most commonly employed [1, 2]. The pulmonary autograft has several advantages including excellent hemodynamics, freedom from anticoagulation therapy, and potential for growth in children. Disadvantages are a two-valve operation, risk of septal ischemia, and pulmonary homograft stenosis and autograft dilation requiring reoperation [1, 3–9]. The pulmonary root, which normally functions within the low pressure pulmonary system, must rapidly adapt to systemic pressure. Ensuing increases in wall stress are associated with remodeling, and autograft dilation can result in aortic insufficiency or aneurysm formation requiring reoperation [2, 4, 7, 9]. These complications have limited widespread application of the procedure.

Comparing native pulmonary and aortic root biomechanics is important for understanding differences in stiffness and wall stress before and immediately after the Ross procedure. Significantly higher wall stress in the autograft as compared with aortic root may lead to pathologic remodeling and dilation; however, biomechanical differences between pulmonary and aortic roots are currently unknown. The objective of this study was to determine and compare material properties of corresponding regions of aortic and pulmonary roots. Biaxial stretch testing was performed within and beyond the physiologic range of strain to develop a constitutive model for future finite element modeling and to assess native pulmonary root mechanics at pulmonary and systemic pressure before remodeling.


   Material and Methods
Top
 Abstract
 Introduction
 Material and Methods
Results
 Comment
 Acknowledgments
 References
 
Fresh porcine hearts (n = 10) from a local abattoir were obtained on the morning of harvest. Porcine tissue was chosen because it has similar anatomy and biomechanics as human tissue and was readily available, unlike human hearts, which were difficult to obtain and test within 24 hours. One human heart, rejected for cardiac transplantation, was obtained after approval by the Institutional Review Board. Five samples were excised from both aortic and pulmonary roots—supravalvular anterior and posterior artery 1 cm distal to sinotubular junction (STJ) and each intrasinus region. Thickness was measured with calipers (Table 1). Tissue from supravalvular ascending aorta (AA) and pulmonary artery (PA) were taken from corresponding locations. Orientation was crucial for determining mechanical properties; thus, care was taken to align specimen in circumferential ({theta}) and longitudinal (L) directions. Samples were washed and stored at 4°C in Dulbecco's saline without calcium and magnesium, and mechanical testing was completed within 24 hours after harvest to preserve tissue mechanical properties.


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  		Table 1 Specimen Dimensions
 
Histology
From an additional four hearts, 10 samples cut as described above were fixed in 10% formalin, embedded in paraffin, and sectioned for histology. Sections were stained with hematoxylin and eosin as well as for collagen and elastin. A cardiovascular pathologist blinded to the specimen region described differences in collagen and elastin.

Biaxial Stretch Testing
A custom-built planar biaxial stretcher was used to determine material properties. Biaxial tensile testing has been previously described [10]. Four 5-0 silk sutures, anchored to each specimen edge using small, barbless fishhooks, were attached to four linear arms of the stretcher. Five black ceramic markers (MO-SCI Corp, Rolla, MO), 250 µm to 355 µm, were placed to create a 3-mm x 3 mm central grid. Tissue was floated in a saline bath at 20°C. Load cells (Model 31/3672-02; Honeywell Sensotec, Columbus, OH), 1,000 g ± 0.1%, on two orthogonal arms were zeroed and monitored to ensure zero force measurement corresponded to resting tissue length.

During extension, data from load cells were amplified (Gould Universal Amplifier, Model 13-4615-58; Gould Instrument Systems, Valley View, OH), 0.01 V, and used to determine force on the sample during deformation. Real-time displacements of marker beads were obtained using a noncontacting CCD camera (30 fps, Model TM9701; Pulnix, Sunnyvale, CA), 0.1 pixels/mm. Images were digitized in MATLAB, version 7.0 (Mathworks, Natick, MA). Marker coordinates were tracked through the loading cycle, and their relative movement was used to calculate Green strains in the principal directions. Strain is the measure of deformation of particles within the material body as they are relatively displaced during soft tissue stretching. As such, strain is dimensionless, which here is expressed as a decimal fraction or percentage as applied to stretching or elongation. Mathematically, Green strain in the circumferential and longitudinal directions was calculated using:


Formula



Formula

where Formula and Formula (Formula ) are stretches in the {theta} and L directions.

Data Collection and Analysis
Samples were tested over a large strain range to compare material properties at pulmonary and systemic pressures immediately after the Ross procedure. Ten preconditioning cycles of 10% stretch, using a triangular waveform at 0.5 Hz were applied. Subsequently, stretching was performed consecutively to 20%, 35%, and 55% peak strain, chosen to approximately correspond to tissue material properties in (1) elastin-dominated, (2) partial fiber engagement or transition, and (3) collagen-dominated regions, respectively [11–13]. Strain at 55% was considered outside physiologic range particularly relevant for pulmonary roots at systemic pressure. Samples were tested using equibiaxial displacement controlled protocols not differential stretch protocols. Differential stretch protocols, in addition to equibiaxial testing, are useful in characterizing coupling between two orthogonal directions [14]. However, we compared stiffness of pulmonary and aortic roots. Since equibiaxial testing is the only protocol that may be used to summarize tissue stiffness differences in two principle strain directions, it was employed here.

For biaxial analysis, we assumed a sample was incompressible—material deformation at constant volume and stress through the tissue thickness was zero. Although arterial tissue is known to have heterogeneous layers and fiber alignment, the tissue was modeled as homogeneous material, and hence global measures of stress and strain were applied [14].

Constitutive Modeling
Stress is defined a measure of the average force acting per unit area of a surface within a deformable body in the deformed configuration, and here is expressed as units of pressure as Pascal (1 Pa is 1 N/m2). Material's response to stress can be described mathematically by a set of constitutive equations, derived from scalar strain energy function W. Mechanical data from supravalvular PA and pulmonary and aortic intrasinuses were fit to a two-dimensional Fung strain energy function, given by


Formula



Formula

where c and Formula are coefficients to the Fung model [10, 12]. Coefficients Formula and Formula are weight fractions of strain components in the principal directions, and Formula represents coupling in two directions [12]. Based on this model, Cauchy stresses for supravalvular PA and pulmonary and aortic intrasinuses were given by


Formula



Formula

Conversely, mechanical data from supravalvular AA samples were fitted to a linear Hookean strain energy function given by Formula


Formula

where coefficients Formula are coefficients to the linear model [10].

Using the Hookean model, Cauchy stresses for supravalvular AA were given by


Formula

A nonlinear regression Levenberg-Marquardt least-squares algorithm in MATLAB was used to fit experimentally obtained stresses to corresponding theoretically calculated stresses. Optimization was iterative and arrived at a solution by minimizing differences between experimentally obtained and calculated Cauchy stresses. Optimization was constrained to produce numerically stable strain energy functions with real world meaning. Coefficient optimization was constrained to ensure convexity and coefficients were solved with a lower boundary of zero [15].

Arterial Stiffness and Dilation
Overall tissue behavior was quantified by tissue stiffness, the numerically determined first derivative of the modeled stress-strain response at a point during deformation. This method allowed for direct comparison of pulmonary and aortic root regions without considering different constitutive equations describing the tissue. Stiffness for each principal direction was calculated at 0.15 strain from 20% stretch, 0.3 strain from 35% stretch, and 0.5 strain from 55% stretch to approximate elastin-dominated, transition, and collagen-dominated regions, respectively [11, 12, 16].

Statistical Analysis
One-way analysis of variance was performed to test for differences in circumferential and longitudinal stiffness between corresponding regions of aortic and pulmonary tissues. Reported values are quoted as mean ± SD, and p values less than 0.05 were considered statistically significant. Statistical analyses were performed using the statistical toolbox in MATLAB.


   Results
Top
 Abstract
 Introduction
 Material and Methods
 Results
Comment
 Acknowledgments
 References
 
Mean thickness of porcine supravalvular AA was significantly greater than that of PA (p = 0.0002), but no significant differences were seen between thickness of aortic and pulmonary sinuses (p = 0.2; Table 1).

Constitutive Modeling
Experimental raw data of representative pig from 55% equibiaxial testing are shown for supravalvular PA and AA, pulmonary and aortic intrasinuses (Fig 1). Both pulmonary and aortic sinuses demonstrated nonlinear material properties with more pronounced nonlinearity than corresponding supravalvular artery. Nevertheless, supravalvular PA was better represented by nonlinear Fung than by linear Hookean strain energy function. As previously reported, the stress-strain response of supravalvular AA was better fit by the linear function even to 0.55 strain [10].


Figure 1
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  		Fig 1. Raw stress-strain data from 55% stretch of representative pig. Supravalvular pulmonary artery (solid triangles) versus ascending aorta (open squares) in (A) circumferential and (B) longitudinal directions; and pulmonary sinuses (solid triangles) versus aortic sinuses (open squares) in (C) circumferential and (D) longitudinal directions.

 
Average coefficients obtained from Levenberg-Marquardt optimization are shown in Table 2. Composite constitutive equations for porcine supravalvular PA and AA and pulmonary and aortic intrasinuses are depicted in Figure 2. Overall, the stress-strain response of supravalvular PA and pulmonary sinuses were qualitatively similar to that of their matched supravalvular AA and aortic sinus counterparts (Fig 2).


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  		Table 2 Coefficients to Fung and Hookean Strain Energy Functions Used to Fit 55% Biaxial Data
 

Figure 2
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  		Fig 2. Modeled stress-strain data derived from averaged coefficients from 55% stretch in (A) circumferential and (B) longitudinal directions. (Ascending aorta = solid circles; aortic sinuses = solid diamonds; pulmonary artery = open circles; pulmonary sinuses = open diamonds.)

 
Arterial Stiffness
In a matched comparison of porcine pulmonary and aortic roots, relative stiffness was dependent on level of strain (Table 3). Strain values of 0.15, 0.30, and 0.5 were tested to represent the elastin-dominated low strain, transition, and collagen-dominated high strain regions, respectively. In the low strain region, supravalvular PA was significantly more compliant than corresponding AA in both circumferential (p < 0.0001) and longitudinal (p < 0.0001) directions. However, this low strain region is likely below physiologic range for both pulmonary and aortic tissue [16, 17]. In the transition region most reflective of physiologic strain, PA remained substantially more compliant circumferentially (p = 0.0005) and longitudinally (p = 0.0002) than AA. At high strain, no significant differences in stiffness were seen between PA and AA circumferentially (p = 0.062); however, PA remained more compliant than AA longitudinally (p = 0.029). This high strain range is pathologic for AA, potentially representing excessive hypertension or aneurysm tissue, whereas for PA, this region likely reflects autograft at systemic pressure before remodeling [16–18].


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  		Table 3 Stiffness of Supravalvular Artery and Intrasinuses Across Strain Range
 
Similar relationships between porcine pulmonary and aortic sinus were not observed (Table 3). No stiffness differences were observed between pulmonary and aortic sinuses circumferentially or longitudinally at low 0.15 or physiologic 0.3 strain (Table 3). It was not until high (0.5) strain that pulmonary sinuses became significantly stiffer than aortic both circumferentially (p = 0.04) and longitudinally (p = 0.013). Figure 3 shows a logarithmic plot of stiffness of each region throughout the entire strain range.


Figure 3
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  		Fig 3. Circumferential stiffness of pulmonary and aortic roots versus strain on a logarithmic scale. (Ascending aorta = solid circles; aortic sinuses = solid diamonds; pulmonary artery = open circles; pulmonary sinuses = open diamonds.)

 
Pulmonary Histology
Pulmonary and aortic root histology stained for elastin (Fig 4) and collagen were compared. Aortic root supravalvular and intrasinus regions had more elastin with denser weave than their pulmonary counterparts. For both pulmonary and aortic roots, supravalvular artery had tight denser weave of elastin than corresponding intrasinuses. Collagen had a more uniform and regular distribution throughout the wall from intima to adventitia in supravalvular artery than intrasinuses. In pulmonary intrasinuses, wider bands of collagen more coarsely distributed were found in the outer one third to one half of the vessel wall near the adventitia as compared with the more tightly woven collagen near the intima.


Figure 4
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  		Fig 4. Elastin-stained histology at low magnification: (A) supravalvular ascending aorta, (B) supravalvular pulmonary artery, (C) aortic intrasinus, and (D) pulmonary intrasinus. Elastin stains black; adventitia on top.

 
Human Sample
Examination of a human tissue sample revealed similar mechanical behavior as porcine tissue. Pulmonary and aortic sinuses had more pronounced nonlinearity than supravalvular PA and AA. At 0.15 strain, human supravalvular PA was more compliant than AA circumferentially (36.59 kPa versus 95.82 kPa, respectively) and longitudinally (91.64 kPa versus 174.77 kPa, respectively). Unlike porcine tissue, by 0.3 strain, PA became stiffer than AA (3,845.34 kPa versus 868.98 kPa circumferential; 9,134.02 versus 1,084.27 kPa longitudinal, respectively) and remained so at 0.5 strain. The transition zones for low, physiologic, and high strain may differ between this human sample and pigs, particularly since the pigs were normal without disease. This human pulmonary sinus was more compliant than aortic sinus at both low (circumferential 184.28 kPa versus 386.28 kPa, respectively; longitudinal 97.99 kPa versus 329.94 kPa, respectively) and higher strain 0.30 (3,978.73 kPa versus 9,618.72kPa circumferential; 487.84 kPa versus 1,1251.07 kPa longitudinal, respectively).


   Comment
Top
 Abstract
 Introduction
 Material and Methods
 Results
 Comment
Acknowledgments
 References
 
In this study, we compared corresponding aortic and pulmonary root material properties demonstrating that aortic and pulmonary sinuses have greater nonlinear stress-strain responses than supravalvular PA and AA. Supravalvular PA was significantly more compliant than AA both circumferentially and longitudinally within the entire strain range, except at high 0.5 strain in the circumferential direction. Conversely, pulmonary sinuses had similar stiffness as aortic through low and transition strain; once high strain was reached, pulmonary sinus stiffness significantly exceeded that of aortic. Histologic differences were also noted, with greater elastin and denser weave in aortic compared with pulmonary counterparts. Supravalvular artery had tighter weave of elastin than intrasinuses for both pulmonary and aortic roots. Differences in collagen distribution were seen with more regular and uniform distribution in supravalvular PA and AA than corresponding intrasinuses, where collagen had wider coarse bands in the outer wall.

Aortic and Pulmonary Root Dynamics
Aortic and pulmonary root dynamics have been examined in vivo [18]. In sheep, sinuses of Valsalva and commissural regions expanded 38% and 64%, respectively, over the cardiac cycle at systemic pressure, whereas STJ and AA expanded 37% and 26%, respectively. In humans, aortic root static pressurization similarly expanded sinuses 20% to 45% with pressures of 60 to 120 mm Hg [19, 20]. Conversely, pulmonary root commissures in sheep expanded 29% versus 28% for STJ and 15% for PA just before the bifurcation at pulmonary pressures. Overall, aortic root expansion at systemic pressure was much greater than pulmonary root expansion at pulmonic pressure. Pulmonary and aortic roots have also been studied within the same pressure ranges using in vitro hydrodynamic studies in pigs [16]. For any given pressure range, PA dilated more than AA at the STJ—from 0 to 30 mm Hg, PA dilated 33% versus 7% for AA; and from 0 to 120 mm Hg, PA dilated 46% versus 35% for AA. Our study corroborates these results: PA was more compliant than AA in the physiologic pulmonary range until the systemic pressure range.

Determining the strain range that corresponds with the pressure range for each type of tissue, pulmonary versus aortic, is difficult. Experimental studies indicate PA nonlinearity to large values of stress may be due to compliance of elastic lamellae and engagement of collagen fibers [13]. In vitro studies suggest that 0.3 to 0.35 strain may be representative of physiologic strain for aortic and pulmonary roots at their respective systemic and pulmonary pressures [16, 17]. When the pulmonary root is exposed initially to systemic pressure (120/80 mm Hg) before remodeling, estimated strain was 46% to 49% with only a 3% strain difference from 80 to 120 mm Hg. We interpreted our 50% strain data to be reflective of PA at systemic pressure [16]. At systemic pressure, PA appeared to have similar stiffness (278 kPa circumferential, 215 kPa longitudinal) at 0.5 strain as AA (291 kPa circumferential, 241 kPa longitudinal) at 0.3 strain. At systemic pressure, pulmonary sinuses were significantly stiffer (1,971 kPa circumferential, 2,055 kPa longitudinal) at 0.5 strain than aortic (276 kPa circumferential, 247 kPa longitudinal) at 0.3 strain. Since PA was significantly more compliant than pulmonary sinuses at 0.5 strain, PA would be expected to initially expand more than pulmonary sinuses based on our results. These native material properties of the pulmonary root acutely dictate how the autograft responds to stress immediately after the Ross operation. However, the impact of these material property differences on biologic remodeling is unknown. Whether the significantly increased stiffness at systemic pressure in the sinuses results in physiologic remodeling versus elastin fiber fragmentation with pathologic weakening and future autograft dilation would be of interest for future study.

Arterial Remodeling
We characterized normal pulmonary and aortic root material properties through a wide strain range, but were limited by not having remodeled pulmonary autografts. Arterial remodeling is an active process that integrates a multitude of stimuli in vivo including blood pressure, flow, and abnormalities in connective tissue with structural and cellular changes in the vasculature [12, 21]. Increases in size and stiffness of arteries are characteristic of hypertensive remodeling due to medial thickening, proliferation of smooth muscle cells, and alteration or damage to extracellular matrix components [21]. After the Ross procedure, rapid transition from pulmonary to systemic pressures can be associated with pathologic autograft remodeling and dilation [1, 2, 4, 6–8, 22].

The majority of Ross patients have some degree of autograft remodeling on echocardiography [6, 7, 22]. Autograft dilation begins within the first week of operation [6]. Annulus and sinus diameters increased immediately after implantation to time of discharge 7 to 10 days postoperatively, followed by significant increases that continued during and beyond the first year. Kouchoukos and colleagues [7] and Luciani and associates [22] demonstrated progressive autograft enlargement over time, most pronounced at the sinus and STJ, with the least at the annulus. Examination of remodeled autograft material properties will be crucial for understanding the link between native material properties and adaptive responses to stress.

To examine autograft structural changes, Carr-White and associates [11] performed uniaxial stretch testing of anterior AA and PA samples taken during the Ross operation before remodeling and one explanted autograft. Their results were similar to ours: PA was highly nonlinear and more compliant than AA. Although a direct comparison to our biaxial stretch data cannot be made, similar trends were observed. At low values of strain, elastic modulus of AA and PA were 130 kPa and 80 kPa, respectively, compared with ours of 113 kPa and 80 kPa. At high strain (approximately 100%), AA and PA elastic modulus were 3,400 kPa and 2,170 kPa, respectively. Throughout the strain range, AA stiffness exceeded that of PA, similar to our observation in porcine AA and PA. Unfortunately, no human intrasinus data were available. Our one human sample seemed more compliant in pulmonary sinus than aortic, similar to porcine data before statistical analyses. Carr-White and colleagues [11] did explant a single autograft within 4 months that appeared stiffer than normal PA, but more compliant than AA, suggesting the autograft gradually evolved to resemble native aorta mechanically. However, one explanted autograft tested uniaxially is insufficient to mechanically describe remodeling [11]. Explanted pulmonary autografts have, however, been studied pathologically, showing dramatic changes in histologic structure [11, 23]. Autograft leaflet structure evolved toward normal aortic valves with trilaminar cuspal structure and collagen architecture and viable valvular interstitial and endothelial cells—all suggestive of repair, adaptation, and growth. In contrast, autograft walls demonstrated extensive structural destruction with granulation tissue and scarring early, and focal loss of normal smooth muscle cells, elastin, and collagen late [23]. It is possible that the significant increases in sinus stiffness we observed at systemic pressure result in elastin fragmentation and structural damage seen eventually in failed autografts. However, our results are based on normal pulmonary roots, not reflective of connective tissue abnormalities. It is unknown whether failed autograft pathologic changes reflect remodeling of structurally normal roots or those with intrinsic abnormalities, namely, bicuspid valve disease.

Clinical Performance
The primary problem with the root replacement method of the Ross procedure is reoperation due to root dilation with aneurysm formation or aortic regurgitation [5, 7, 8, 22]. Autograft dilation has been described occurring in the sinuses or STJ alone or in both. These dilation modes seen by magnetic resonance imaging may be dependent on the proportion of PA distal to STJ incorporated in the autograft as well as surgical implantation technique [22, 24, 25]. Since pulmonary sinus is substantially stiffer than the supravalvular artery at high strain/systemic pressure, our results suggest a greater expansion of PA than sinus initially. If these differences were sustained after remodeling, then the length of PA incorporated in the autograft should be minimized, and reinforcement of the STJ considered. However, we also demonstrate significantly higher stiffness in pulmonary versus aortic sinuses initially at high strain that may contribute to structural damage, in which case, sinus reinforcement may also be recommended. Chronically, autograft sinuses can dilate, leading to reoperation, and no definitive answer regarding the optimal Ross implantation technique is available based on current clinical outcomes [1–3].

Comprehensive long-term follow-up data after the Ross operation is sparse and highly variable. Attempts have been made to support the autograft using the root inclusion technique, which maintains the native aortic walls, or by applying a Dacron jacket around the autograft [26]. Some surgeons have advised against root replacement, citing better patient outcomes with subcoronary implantation [1, 3]. In addition to retrospective clinical studies, experimental studies such as this one and future experimental and computational studies of remodeled autografts may shed light on autograft adaptive responses to systemic blood pressure. Such data may aid the determination of optimal implantation techniques or choices of mechanical support to prevent autograft failure.


   Acknowledgments
Top
 Abstract
 Introduction
 Material and Methods
 Results
 Comment
 Acknowledgments
References
 
We thank Dr Philip Ursell for his histologic characterization of collagen and elastin distribution and the California Transplant Donor Network for the rejected human heart.


   References
Top
 Abstract
 Introduction
 Material and Methods
 Results
 Comment
 Acknowledgments
 References
 

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