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J Thorac Cardiovasc Surg 2003;125:1405-1411
© 2003 The American Association for Thoracic Surgery

Surgery for Acquired Cardiovascular Disease

A new equine pericardial stentless valve

From the Department of Cardio-vascular Surgery, Centre Hospitalier Universitaire de Sherbrooke, Sherbrooke, Quebec, Canada,^a and the Department of Cardio-vascular Surgery, Centre Hospitalier Universitaire Vaudois (CHUV), Lausanne, Switzerland.^b

Received for publication May 30, 2002. Revisions requested Aug 26, 2002; revisions received Sept 9, 2002. Accepted for publication Oct 25, 2002. Address for reprints: Xavier Mueller, MD, Department of Cardio-vascular Surgery, Centre Hospitalier Universitaire de Sherbrooke, 3001, 12^e Avenue Nord, Sherbrooke, Quebec, Canada J1H 5N4 (E-mail: xavier.mueller{at}usherbrooke.ca).

	Abstract

Top Abstract Introduction Material and methods Results Discussion Appendix: Discussion References

 
Objective: Small aortic valve replacement remains a challenging hemodynamic problem. A new bioprosthesis (3F Therapeutics, Lake Forest, Calif) was designed to further improve the hemodynamic performance currently achieved with stentless bioprostheses. This valve consists of a tubular structure assembled from 3 equal sections of equine pericardial material, with virtually no foreign material except for a thin polyester ring. Its hemodynamic performance was compared with that of a commercially available stentless prosthesis in a bovine model.
Patients and Methods: Twelve calves (55 ± 2.8 kg) received a 19-mm 3F valve (3F group, n = 6) or a 19-mm stentless control valve (control group, n = 6). The animals were fully equipped for hemodynamic monitoring and transvalvular gradient measurements. After implantation, dopamine was infused in increasing doses, and the hemodynamic values were recorded at each step of 100-µg/min increase. Linear regression analysis was applied for group comparison of each variable.
Results: The mean transvalvular gradient at 4.5 L/min was 3.48 ± 0.14 mm Hg (95% confidence interval) in the 3F group and 5.72 ± 0.28 mm Hg in the control group (P < .0001) and at 6.5 L/min, 7.4 ± 1.55 mm Hg, and 11.13 ± 0.18 mm Hg, respectively (P < .0001). The effective orifice area at 4.5 L/min was 2.4 ± 0.03 cm² in the 3F group and 1.86 ± 0.02 cm² in the control group (P < .0001) and at 6.5 L/min, 2.41 ± 0.04 cm², and 1.96 ± 0.02 cm², respectively (P < .0001).
Conclusions: This new bioprosthesis without a stent and without a supporting wall that has its commissures fixed directly to the aorta outperforms in vivo standard stentless prostheses in the immediate postimplant period.

	Introduction

Top Abstract Introduction Material and methods Results Discussion Appendix: Discussion References

 
Although stented bioprostheses avoid the need for anticoagulation, valve stents have 2 main drawbacks. First, stents have been shown to increase stress on the biologic cusp tissue in fatigue-testing models. ¹ This feature has been involved as the source of leaflet calcification, with subsequent limited valve durability. Second, the obstructive nature of rigid stents leads to a nonphysiologic flow pattern and residual gradient, which would inhibit complete resolution of left ventricular hypertrophy and lead to lower long-term survival. ²

Stentless aortic xenograft valves offer the potential to overcome the problems associated with stented valves and to mimic the hemodynamic performance of homografts. ³ Despite excellent initial results after its clinical introduction, ^4,5 premature structural deterioration because of poor preservation methods hampered its further application. The introduction of glutaraldehyde by Carpentier and associates ⁶ in 1968, with the subsequent widespread use of stented bioprostheses in the early 1980s, renewed the interest in stentless xenograft valves. David and colleagues ⁷ began to implant stentless porcine aortic valves in sheep in 1988, with encouraging hemodynamic results. The 1990s have witnessed the introduction of innovative stentless aortic valves, among which the United States Food and Drug Administration approved the Toronto SPV valve (St Jude Medical, Inc, St Paul, Minn) and the Freestyle valve (Medtronic, Inc, Minneapolis, Minn).

Nevertheless, any aortic valve substitute, even the stentless valve, has its inherent obstructive feature. In both the Toronto SPV and Freestyle valves, the leaflets are supported by a Dacron cloth, as well as by the aortic wall of the xenograft. This support apparatus represents additional material within the recipient root. A new bioprosthesis (3F Therapeutics, Lake Forest, Calif) was designed to further diminish these obstructive features. This valve consists of a tubular structure assembled from 3 equal sections of equine pericardial material with virtually no foreign material except for a thin polyester ring. Because residual gradient is a concern, especially in patients with small aortic roots, this experimental study was designed to compare the hemodynamic performance of the 19-mm 3F valve with that of the Toronto SPV valve of corresponding size in a calf model.

	Material and methods

Top Abstract Introduction Material and methods Results Discussion Appendix: Discussion References

 
Study design
The protocol described herein was reviewed and approved by the Committee on Animal Care, Office Vétérinaire Cantonal, Lausanne, Switzerland. All animals received care in compliance with the "Principles of Laboratory Animal Care" formulated by the National Society for Medical Research, and the "Guide for the Care and Use of Laboratory Animals" prepared by the Institute of Laboratory Animal Resources, National Research Council, and published by the National Academy Press, revised 1996.

Valve design
The 3F valve is a low-pressure glutaraldehyde-fixed pericardial bioprosthesis that consists of a tubular structure assembled from 3 equal sections of equine pericardial material (Figure 1). The 3 leaflets have been assembled with locking sutures. Three tabs have been placed at the outflow, simulating commissures that allow fixation in the vicinity of the sinotubular junction. These tabs are reinforced with equal sections of polyester material that also will serve to promote ingrowth of the tab into the aortic wall. Leaflets and tabs are one integral structure. At the inflow area of the tubular structure, a slightly scalloped polyester ring has been incorporated to allow easy suturing of the bioprosthesis to the bottom of the aortic sinus, thus securing it and minimizing the potential for perivalvular leakage. It is expected that this ring will help the promotion of ingrowth of fibrous material to anchor the valve firmly to the root of the aorta.

View larger version (129K): [in this window] [in a new window]   Fig. 1. 3F aortic pericardial bioprosthesis.

View larger version (91K): [in this window] [in a new window]   Fig. 2. A, Primal tubular structure of human heart valve during development in utero. B, The white dots indicate the points of fixation of the tubular structure (ie, the commissural areas). When fluid flows through this tubular structure, the points of fixation will automatically determine the form of the valve.

Because of the short length of the ascending aorta, both the right carotid artery and the aortic arch were cannulated. For the venous return, a double-stage cannula was introduced into the right atrium. A standard cardiopulmonary bypass (CPB) circuit was connected. Whole-body perfusion was then instituted, with the blood flow rate maintained by a roller pump at 60 to 80 mL · kg^-1 · min^-1. Body temperature was lowered to 28°C. After the brachiocephalic trunk and the ascending arch had been crossclamped with a single clamp, 1 L of cold crystalloid solution was injected through a cardioplegia needle at the base of the aorta. The distal ascending aorta was transected to expose the aortic valve and to allow maximal length of the aortic root for proper implantation of the valve. The transection provides better exposure than when the aorta is not completely transected, and it preserves the normal anatomic structure of the root. The native aortic cusps were then carefully excised.

The 12 animals were randomly assigned to either a 3F valve (3F group) or a Toronto SPV valve (control group). Before implantation, the aortic annulus size was assessed with a 19-mm Hegar dilator. When a 3F valve was inserted, its polyester ring was secured to the annulus with 2-0 Ti-Cron (United States Surgical Corp, Norwalk, Conn) interrupted sutures. Then each tab was fixed to the aortic wall with a transmural 4-0 Prolene (Ethicon, Inc, Somerville, NJ) U stitch.

When a Toronto SPV valve was used, its annulus was secured in a similar fashion. The valve was oriented such that the largest opening faced the left ostium. After tying the proximal suture line, the commissural posts were held in place with stay sutures of 4-0 Prolene to achieve proper orientation. Then the distal suture line was performed with a running 4-0 Prolene suture.

After valve implantation, the aorta was partially closed with a running 4-0 Prolene suture. Once the deairing maneuver was performed, the aortotomy was sealed and the crossclamp was removed. After rewarming, the heart was defibrillated if necessary. When normal sinus rhythm was established, the animal was carefully weaned from CPB with a dopamine support of 200 µg/min, and volume was adjusted as needed.

Hemodynamic recording
After weaning from CPB, a period of 20 minutes of stabilization was allowed before recording the cardiac output, the mean transvalvular gradient, and the maximal transvalvular gradient. Data were recorded digitally online with a sampling rate of 250 Hz per channel (Transceiver unit TRx 001; Sonometrics, London, Ontario, Canada). Mathematic analysis of the data was performed offline with a software package for cardiovascular analysis (Sonosoft version 3.1.3., Sonometrics). The hemodynamic data were obtained by measuring 5 consecutive beats twice at 1-minute intervals. Then the dose of dopamine was increased at 20-minute intervals by using 100 µg/min. At the end of each period, the hemodynamic data were recorded. The procedure was stopped when the heart rate exceeded 120 beats/min.

Calculations
Valve orifice area (VOA) in square centimeters was calculated by the formula of Gorlin and Gorlin ⁸:
VOA = CO/44.3 (SEP) (HR) MPG
where CO is the cardiac output (in liters per minute), SEP is the systolic ejection period (in milliseconds), HR is the heart rate (in beats per minute), and MPG is the mean pressure gradient (in milliliters of mercury).

The valve resistance (VR; dyne · cm · s^-5) was calculated according to the following formula ⁹:
VR = 1.33 (MPG) (HR) (SEP)/CO.

Statistics
Values are given as means ± SD. Correlations between hemodynamic parameters (mean transvalvular gradient, maximal transvalvular gradient, valve orifice area, and valve resistance) and cardiac output were performed with linear regression analysis, according to the least-squares method. Then each hemodynamic parameter, as determined by means of regression analysis, was further examined with the unpaired Student t test between the 2 valve groups.

	Results

Top Abstract Introduction Material and methods Results Discussion Appendix: Discussion References

 
The 3F aortic valve could be implanted in all the animals, with successful weaning from CPB and complete hemodynamic measurements obtained. Sinus rhythm was maintained during the complete set of measurements. Four steps of measurements could be performed in each animal, except for one animal in the standard group, in which 5 steps could be performed. Therefore, for each parameter, 240 measurements were available in the 3F group and 250 measurements in the standard group.

The mean heart rate was 89.7 ± 11.9 beats/min during the first series of measurements and 116.4 ± 3 beats/min during the last series. Aortic crossclamp time was 74 ± 7 minutes for the 3F group and 92 ± 5 minutes for the control group (P < .0001).

Linear regression analysis of the mean transvalvular gradient according to the cardiac output is shown on Figure 3. The difference between both groups is statistically significant (P < .0001), with gradients of 3.48 ± 0.14 mm Hg (95% confidence interval [CI]) for the 3F group versus 5.72 ± 0.28 mm Hg for the standard group at 4.5 L/min and gradients of 7.4 ± 1.55 mm Hg versus 11.13 ± 0.18 mm Hg at 6.5 L/min, respectively.

View larger version (12K): [in this window] [in a new window]   Fig. 3. Mean transvalvular gradient: regression analysis according to the cardiac output (CO). Dotted lines, 95% CI.

View larger version (12K): [in this window] [in a new window]   Fig. 4. Maximal transvalvular gradient: regression analysis according to the cardiac output (CO). Dotted lines, 95% CI.

View larger version (12K): [in this window] [in a new window]   Fig. 5. Valve orifice area (VOA): regression analysis according to the cardiac output (CO). Dotted lines, 95% CI.

View larger version (12K): [in this window] [in a new window]   Fig. 6. Valve resistance (VR): regression analysis according to the cardiac output (CO). Dotted lines, 95% CI.

	Discussion

Top Abstract Introduction Material and methods Results Discussion Appendix: Discussion References

The principle of the 3F valve reverses the usual concept of imitating the final form of a defined structure. On the contrary, the 3F design is derived from the primal tubular structure of human heart valves found during their original development. The 3 points of fixation of this tubular structure will determine the subsequent form of the valve once blood flow is established (ie, once its function is acquired). This concept allows for a dramatic simplification of the prosthetic structure. When compared with the Toronto SPV valve and the Freestyle valve, 3 features have been fundamentally modified: the absence of a supporting xenograft aortic wall, the reduction of synthetic cloth, and the absence of a muscle bar.

First, there is no supporting wall. This might reduce transprosthetic gradients in 2 ways. On the one hand, the xenograft aortic wall itself is an obstructive element once implanted within the native aortic root. On the other hand, analysis of hemodynamics and left ventricular mass regression after implantation of the Toronto SPV valve, ^10-12 as well as the Freestyle valve, ^13-15 have shown repeatedly a significant decrease in transvalvular gradient over the first year after implantation. Although ventricular remodeling has been involved as the major contributor to this decrease beyond the first 3 to 6 postoperative months, absorption of the hematoma in the potential space between the porcine and the native aortic walls appears more likely to be the predominant mechanism in the early postoperative period. The absence of any supporting structure offers to the 3F valve the potential for optimal gradient as soon as the implantation time. This hypothesis is supported by the analysis of gradient profile over the time of the Freestyle valve implanted as a root replacement. ¹⁶ In contrast to the subcoronary implant technique and the inclusion technique, early and late gradients were both low and did not change with time.

Second, the absence of foreign material except for a thin polyester ring contrasts with the heavy external Dacron support found especially in the Toronto SPV valve. This avoids any unnecessary increase in thickness of the xenograft. Therefore, artificially induced stiffness or stress on the tissue can be reduced. In keeping with this concept, the reduction of synthetic supporting structure allows for the reduction of suture lines.

Third, the pericardial structure of the 3F valve avoids the presence of the unfavorable muscle bar characteristic of the base of the porcine right coronary leaflet, therefore maximizing the valve orifice area. Moreover, coronary ostial position relative to the annulus might further exaggerate the problem of the muscle bar. Westaby and coworkers ¹³ have described the potential pitfall of Freestyle valve linked with the Dacron structure, particularly over the muscle bar area, thus generating increased transvalvular gradients.

One of the primary aims of eliminating stents from bioprosthetic valves was to use the sinus of Valsalva of the native or xenograft aortic root as a functional stent. This would allow the leaflets to open more fully as commissures are pulled apart during systole. Despite the absence of rigid stents, all currently available stentless valves have some xenograft or synthetic supporting structure, which might be viewed as soft stents. The 3F valve explores the stentless concept further by eliminating supporting components of any kind.

A substantial advantage of 3F valve implantation lies in its shorter time of implantation. Instead of 2 suture lines, the technique involves only one inflow suture line and 3 commissural stitches, allowing for reduced crossclamp time. Moreover, the pliability of the valve facilitates its manipulation, which might be particularly relevant in a small aortic root. The stability of the 3F valve orifice area at all cardiac output values tested might be related to its simplified and pliable structure, contrasting with the greater inertia of the Toronto SPV valve, especially at low flow rates.

Several features of the study deserve comment. A calf model has several advantages over other animal models for acute hemodynamic study of aortic prosthetic valves. A pig model is limited by post-CPB complications, sheep have a narrow and fragile aortic root, and goats have limited availability.

An in vivo rather than in vitro study was performed because in vitro values have been shown to be unable to predict the behavior of stentless prostheses in vivo. ¹⁷ Valve orifice areas were found to be lower in vivo because stentless valves somewhat collapse when inserted in vivo, whereas in vitro the prostheses are fully expanded and behave more like stented prostheses.

The 19-mm size was chosen for 2 reasons. On the one hand, the performance of a valve is better challenged with small-sized versions. On the other hand, a small aortic root is an increasing problem with the aged population. In a multicenter study, Hvass and associates ¹⁸ found 171 (30%) of 561 patients with a small aortic root as defined by a 19- to 21-mm aortic annulus. In patients older than 80 years of age, the percentage reached 40%.

A protocol testing hemodynamics over a wide range of cardiac output allows a better evaluation of valve performance when compared with most clinical studies performing echocardiographic evaluation at rest only. Moreover, most of these echocardiographic studies do not mention the cardiac output value at the time of the examination. ^10,11,13-15 This lack makes any comparison difficult because mean transvalvular gradient, maximal transvalvular gradient, and valve orifice area have all been demonstrated to be flow dependent. ^19-21

A comparative study was designed because absolute values found in the experimental setting cannot be extrapolated directly to human subjects, mainly because of anatomic differences. For instance, in this study maximal transvalvular gradients have been found to be disproportionately high in comparison with mean transvalvular gradients. This is likely because of the particularly thickened aortic wall of the calf, which offers a low compliant outflow chamber to transvalvular blood flow.

As for limitations, this acute experimental model does not evaluate the long-term outcome of prosthetic valves, namely degeneration and durability. The theoretical reduction in leaflet stress leading to improved durability remains a challenging issue that needs further evaluation. Second, the sinotubular junction size in human subjects might be larger than that of the annulus. In the Toronto valve experience, David ²² recommended that the prosthesis should be sized at the sinotubular junction. This usually leads to an oversizing of the prosthesis when referred to the annulus. The problem of 3F valve sizing deserves further evaluation. Finally, candidates for aortic valve replacement might exhibit a dilative tendency at the level of the sinotubular junction, which might develop over time. This issue has been shown to have relevant consequences in terms of late prosthetic regurgitation. ²³ Whether the recommendation of banding the sinotubular junction at the time of Toronto SPV valve implantation should be applied to the 3F valve needs to be evaluated.

	Appendix: Discussion

Top Abstract Introduction Material and methods Results Discussion Appendix: Discussion References

 
Dr Henry M. Spotnitz (New York, NY). There are other issues about valve design that you did not address in your presentation. Can you tell us how the pericardium was preserved if at all, whether you have any information about long-term durability, issues of valvular regurgitation, and how critical you consider the fixation of those tags to the aortic wall? Is that a critical issue, or can they be sewn anyplace? Does it affect the amount of regurgitation? What else can you tell us about that?

Dr Mueller. I will start first with the last question. The very critical point, as for the other stentless valve, is the anchoring of these 3 points. They have to be very precisely positioned. They should be fixed at about the level of the sinotubular junction or just below it. If you do not fix them properly, and if the tubular form is not achied properly—you should have enough tension between the inflow of the valve and the outflow—you will not achieve such a gradient.

Fixation is standard low-pressure glutaraldehyde fixation.

Long-term durability has been tested in vitro according to the ISO standards, and no specific defect has been noted.

Dr Wolfgang Goetz (Regensburg, Germany). This valve looks very nice in the normal aortic root, but what if this root is dilated? I would expect some tethering on the commissures, which are fixed on only 3 points, and if there is a dilatation, I would expect some central regurgitation.

Dr Mueller. Of course. This point has been raised recently with the Toronto valve by Tirone David. His group observed some aortic regurgitation with dilatation of the sinotubular junction, and they suggested wrapping this area with a tape to fix it or to reduce it. The present experimental study did not address this point, but this finding should be valid for this stentless valve as well.

Dr Goetz. In contrast to a stentless valve, the commissures are fixed on only one point. Therefore, if there is any kind of dilatation, you might have a huge tethering there, and it might rupture at these commissures.

Dr Mueller. There are 2 possibilities to address this problem. First, you might oversize the valve. Second, you might reduce the sinotubular junction, fix it, or both. This is obviously an important point.

Dr José M. Revuelta (Santander, Spain). We just started using this valve in our institution as part of an international multicenter study, and we found it easy to implant. It is critical to produce a good attachment at the top.

Concerning the previous question, if the aorta is dilated and the patient has a dilated aorta, we need to remodel the sinotubular junction. I think that this is not a contraindication.

My question is this: Did you use, in your clinical case, Prolene running sutures or interrupted sutures? We prefer 4-0 Ethibond interrupted sutures because it is only a 1-line suture. If something happened, the patient would have a serious problem, particularly with a fragile and calcified annulus, as is true in most of the cases.

Dr Mueller. We used interrupted sutures in the clinical case. In the experimental setting we did the same because of the small size of the annulus and because we did not want to have any purse-string effect.

Dr Juan Carlos Chachques (Paris, France). First, I want to ask why you use equine pericardium. There are some differences with bovine histologically, because I think it would be more difficult to use this material. Second, you are cutting 3 pieces of your valve, but you make a tube. Why don't you make only one suture instead of taking 3 parts to make your valve?

Dr Mueller. In answer to the first question, equine pericardium is thinner and more pliable. That was the idea behind that. Now, for the tubular structure, if you take 3 equal pieces of pericardium that are sutured together, you will already put the final form in line when the distal end of the 3 linear sutures are fixed directly to the aorta. In this way you obtain spontaneously a trileaflet structure once it is closed by the backflow.

Dr Eugene A. Grossi (New York, NY). I enjoyed the presentation. I especially enjoyed the picture of the Hegar dilators in showing the difference in internal diameters of the valve.

Are you postulating that there is a difference in hemodynamic performance because one valve has an internal diameter of 17 mm versus a 19-mm internal diameter?

What is the method of sizing the valves? Because different manufacturers use various methods for determining a particular size, can this account for the hemodynamic differences,r do you believe there are intrinsic differences in the valvular function other than size?

Dr Mueller. Yes, the size was one of the main issues, especially with the Toronto valve, which has been recently compared with the Carpentier-Edwards valve in that respect. With the 3F valve, the 19-mm size corresponds to the inner diameter. With the 19-mm Hegar dilator, the outer diameter of the 3F valve is 20 mm. Therefore, the wall of this valve is very thin.

The effect of the size characteristics of the valve on the hemodynamics is definitely important, particularly for a small-sized valve at low cardiac output. At higher output, you might expect that the flow will push all the structures toward the recipient aortic wall, but at lower output, these structures might partly obstruct the outflow tract, especially with a small-sized valve.

	Acknowledgments

 
We thank Philippe Frascarolo, PhD, for his help in the statistical evaluation of this article.

	Footnotes

 
Read at the Eighty-second Annual Meeting of The American Association for Thoracic Surgery, Washington, DC, May 5-8, 2002.

	References

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