There has been renewed interest in developing polymer valves. Here, the objectives are to demonstrate the feasibility of modifying a commercial pulse duplicator to accommodate tri-leaflet geometries and to define a protocol to present polymer valve hydrodynamic data in comparison to native and prosthetic valve data collected under near-identical conditions.
Limitations of currently available prosthetic valves, xenografts, and homografts have prompted a recent resurgence of developments in the area of tri-leaflet polymer valve prostheses. However, identification of a protocol for initial assessment of polymer valve hydrodynamic functionality is paramount during the early stages of the design process. Traditional in vitro pulse duplicator systems are not configured to accommodate flexible tri-leaflet materials; in addition, assessment of polymer valve functionality needs to be made in a relative context to native and prosthetic heart valves under identical test conditions so that variability in measurements from different instruments can be avoided. Accordingly, we conducted hydrodynamic assessment of i) native (n = 4, mean diameter, D = 20 mm), ii) bi-leaflet mechanical (n= 2, D = 23 mm) and iii) polymer valves (n = 5, D = 22 mm) via the use of a commercially available pulse duplicator system (ViVitro Labs Inc, Victoria, BC) that was modified to accommodate tri-leaflet valve geometries. Tri-leaflet silicone valves developed at the University of Florida comprised the polymer valve group. A mixture in the ratio of 35:65 glycerin to water was used to mimic blood physical properties. Instantaneous flow rate was measured at the interface of the left ventricle and aortic units while pressure was recorded at the ventricular and aortic positions. Bi-leaflet and native valve data from the literature was used to validate flow and pressure readings. The following hydrodynamic metrics were reported: forward flow pressure drop, aortic root mean square forward flow rate, aortic closing, leakage and regurgitant volume, transaortic closing, leakage, and total energy losses. Representative results indicated that hydrodynamic metrics from the three valve groups could be successfully obtained by incorporating a custom-built assembly into a commercially available pulse duplicator system and subsequently, objectively compared to provide insights on functional aspects of polymer valve design.
Heart valve disease often results from degenerative valve calcification1, rheumatic fever2, endocarditis3,4 or congenital birth defects. When valve damage occurs, causing stenosis and/or regurgitation valve prolapse and cannot be surgically repaired, the native valve is usually replaced by a prosthetic valve. Currently available options include mechanical valves (cage-ball valves, tilting disk valves, etc.), homograft, and bioprosthetic valves (porcine and bovine valves). Mechanical valves are often recommended for younger patients based on their durability; however the patient is required to remain on anticoagulant therapy to prevent thrombotic complications5. Homograft and biological prosthetic valves have been effective choices to avoid blood thinner therapy; however, these valves have elevated risk for fibrosis, calcification, degeneration, and immunogenic complications leading to valve failure6. Tissue-engineered valves are being investigated as an emerging technology7-9, but much still remains to be uncovered. Alternative durable, biocompatible, prosthetic valves are needed to improve the quality of life of the heart valve disease patients. Again, this valve design could replace the bioprosthesis used in transcatheter valve technology, with transcatheter approaches showing the potential for transforming the treatment of selected patients with heart valve disease10.
As stated by current standards, a successful heart valve substitute should have the following performance characteristics: "1) allows forward flow with acceptably small mean pressure difference drop; 2) prevents retrograde flow with acceptably small regurgitation; 3) resists embolization; 4) resists hemolysis; 5) resists thrombus formation; 6) is biocompatible; 7) is compatible with in vivo diagnostic techniques; 8) is deliverable and implantable in the target population; 9) remains fixed once placed; 10) has an acceptable noise level; 11) has reproducible function; 12) maintains its functionality for a reasonable lifetime, consistent with its generic class; 13) maintains its functionality and sterility for a reasonable shelf life prior to implantation."11. Some of the shortcomings of existing valve prostheses may potentially be overcome by a polymer valve. Biocompatible polymers have been considered top candidates based on biostability, anti-hydrolysis, anti-oxidation, and advantageous mechanical properties such as high strength and viscoelasticity. In particular, elastomeric polymers may provide material deformation resembling native valve dynamics. Elastomers can be tailored to mimic soft tissue properties, and they may be the only artificial materials available that are bio-tolerant and that can withstand the coupled, in vivo, fluid-induced, flexural and tensile stresses, yet, move in a manner resembling healthy, native valve motion. Moreover, elastomers can be mass-produced in a variety of sizes, stored with ease, are expected to be cost-effective devices and can be structurally augmented with fibrous reinforcement.
The concept of the use of polymer materials to assemble a tri-leaflet valve is not new and has been the subject of several research investigations over the last 50 years12, which were abandoned largely due to limited valve durability. However, with the advent of novel manufacturing methodologies13,14, the reinforcement of polymer materials15,16 and potentially seamless integration of polymer valve substitutes with transcatheter valve technology, there has recently been a renewed interest and activity in developing polymer valves as a potentially viable alternative to currently available commercial valves. In this light, a protocol for enabling testing of these valves to assess hydrodynamic functionality is the first step in the evaluation process; yet commercially available pulse simulator systems generally do not come equipped to accommodate tri-leaflet valve designs and contain an annular spacing to insert commercially available heart valves (e.g. tilting disc, bi-leaflet mechanical heart valves). Secondly, polymer valves are an emerging technology whose hydrodynamics can only be assessed in a relative context. Even though native heart valve pressure and flow data is available, it is important to conduct testing of native aortic porcine valves, which are biologically similar to human valves, using the same pulsatile simulator that is used to evaluate the polymer valves so as to account for measurement differences that may be system dependent. Thus, the goal of this study was to demonstrate how a commercially available pulse simulator can be fitted with an assembly to accommodate tri-leaflet valve constructs and to systematically evaluate polymer valve hydrodynamic metrics in a relative context in comparison to mechanical and native porcine heart valve counterparts. In our case, novel tri-leaflet silicone polymer valves previously developed at the University of Florida 13 comprised the polymer valve group.
1. Preparation
2. Native Aortic Valve Dissection
3. Polymer and Native Valve Suturing Process
4. Hydrodynamic Evaluation
Note: Actual protocol will vary depending on specific pulse duplicator system being used. All information caontained herein used the ViVitro Pulse Duplicator Sysytem (ViVitro Labs, Inc., Vancouver, BC).
5. Post Processing
Representative flow and pressure waveforms are shown in Figures 3, 4 and 5. The plots were averaged over the sample size of valves tested for each group, which was, n = 5, 4, and 2 valves for polymer, native porcine and bi-leaflet groups, respectively. The mean hydrodynamic metrics and the standard error of the mean for these sample sizes are presented in Table 1.
Figure 1. (a) Schematic of the ViVitro pulse duplicator system showing the primary components that implement a Windkessel model for physiologically relevant flows (figure presented here with permission from ViVitro Systems, Inc, BC, Canada). (b) Rapid prototyped valve holder configuration to suture and secure silicone or native porcine valves in-place. (c) Modification of the ViVitro pulsatile loop to accommodate tri-leaflet valve constructs. Click here to view larger figure.
Figure 2. (a) Native porcine valve. (b) Top view of polymer valve leaflets. (c) Side view of polymer valve after suturing and securing in-place within valve-holder. (d) Saint Jude bi-leaflet mechanical valve. Click here to view larger figure.
Figure 3. Mean instantaneous flow rates of the 3 valves tested (n = 5, 4, and 2 valves for polymer, native porcine and bi-leaflet, respectively). Flow rate was measured using an electromagnetic flow meter connected to a noninvasive flow probe placed at the interface location of the ventricle and aortic chambers (see Figure 1a). Click here to view larger figure.
Figure 4. Mean instantaneous ventricular pressure of the 3 valves tested (n = 5, 4, and 2 valves for polymer, native porcine and bi-leaflet, respectively). Ventricular pressure was measured in the ventricle chamber using a micro-tip pressure transducer. Superimposed literature ventricular pressure values for native and bi-leaflet valves (Diameter: 29 mm) were obtained from18 and19, respectively. Click here to view larger figure.
Figure 5. Mean instantaneous aortic pressure of the 3 valves tested (n = 5, 4, and 2 valves for polymer, native porcine and bi-leaflet, respectively). Aortic pressure was measured just downstream from the aortic valve position using a micro-tip pressure transducer. Superimposed literature aortic pressure literature values for native and bi-leaflet (Diameter: 29 mm) valves were obtained from18 and19, respectively. Click here to view larger figure.
Bi-leaflet (n=2) | (Polymer n=5) | Porcine (n=4) | ||||
Data Description | Mean | SEM | Mean | SEM | Mean | SEM |
Aortic Orifice Area [P] (cm2) | 3.143 | 2.697 | 2.920 | 1.306 | 2.516 | 1.258 |
Aortic Orifice Area [F] (cm2) | 7.940 | 1.286 | 4.613 | 2.063 | 3.975 | 1.988 |
Aortic Orifice Area [H] (cm2) | 7.516 | 1.633 | 4.575 | 2.046 | 3.942 | 1.971 |
Forward Flow Pressure Drop [P] (mmHg) | 17.000 | 0.054 | 22.284 | 12.007 | 40.795 | 11.670 |
Forward Flow Pressure Drop [F] (mmHg) | 0.410 | 0.210 | 30.424 | 9.235 | 29.766 | 9.733 |
Forward Flow Pressure Drop [H] (mmHg) | 26.520 | 0.120 | 50.790 | 4.230 | 5.610 | 4.970 |
Trans-Aortic Max Pressure (mmHg) | 15.850 | 12.400 | 60.930 | 20.470 | 75.250 | 17.470 |
Aortic RMS Forward Flow Rate [P] (ml/sec) | 88.280 | 11.110 | 162.120 | 24.970 | 189.080 | 32.610 |
Aortic RMS Forward Flow Rate [F] (ml/sec) | 193.570 | 3.820 | 204.560 | 6.680 | 177.310 | 2.630 |
Aortic RMS Forward Flow Rate [H] (ml/sec) | 197.790 | 0.630 | 174.760 | 11.530 | 182.680 | 3.160 |
Aortic Forward Volume (ml) | 68.180 | 6.430 | 55.390 | 3.660 | 64.200 | 1.750 |
Aortic Closing Volume (ml) | 62.260 | 0.860 | 32.990 | 9.820 | 45.260 | 11.990 |
Aortic Leakage Volume (ml) | 60.140 | 3.470 | 33.090 | 9.220 | 56.130 | 11.260 |
Total Regurgitant Volume (ml) | 122.400 | 4.320 | 66.080 | 17.200 | 101.390 | 23.160 |
TransAortic Forward Flow Energy Loss (mJ) | 80.321 | 4.65 | 115.287 | 17.354 | 184.325 | 12.354 |
TransAortic Closing Energy Loss (mJ) | 25.231 | 0.589 | 29.52 | 6.872 | 12.354 | 4.874 |
TransAortic Leakage Energy Loss (mJ) | 87.219 | 13.242 | 84.02 | 12.205 | 97.029 | 25.047 |
TransAortic Total Energy Loss (mJ) | 192.771 | 23.51 | 228.827 | 47.254 | 293.708 | 36.483 |
Table 1. Mean and Standard Error of the Mean (SEM) Hydrodynamic metrics computed for the heart valves tested (n = 5, 4, and 2 valves for polymer, native porcine and bi-leaflet, respectively). The following intervals should be noted: P: interval that starts and ends with 0 TVP, F: interval with forward flow and H: interval starting with 0 TVP and ending with 0 flow. Mean diameters of the valves were as follows: Polymer valve (n=5): 22 mm; Native porcine valve (n=4): 20 mm; bi-leaflet (n=2): 23 mm. Small sample size for bi-leaflet valve was due to limited samples available for research use; the two bi-leaflet valves tested were previously donated to the Biomedical Engineering Department at Florida International University by Saint Jude Medical (Saint Paul, MN).
In this study, we have demonstrated the utility of modifying a commercially available pulsatile duplicator unit to accommodate tri-leaflet valve geometries so that hydrodynamic testing of polymer and native porcine valves can be performed. Specifically in our case, the system modified was a ViVitro left heart and systemic simulator system (Figure 1a) controlled via the ViViTest data acquisition system (ViVitro Systems, Inc, Victoria, BC, Canada). However, the system is not unlike several in vitro, pulsatile flow loops which all utilize a two-component Windkessel model to mimic flow and pressure waveforms of relevance to the human circulation22-25. These two-component Windkessel systems typically consist of a pulsatile pump, a compliance chamber that mimics the distensibility of the arteries, and a peripheral resistance controller that can be used to regulate the vascular resistance. The equation that describes the two-component Windkessel model is:
where C is the compliance, R the resistance, Q(t) is the volumetric flow rate as a function of time and P is the arterial pressure (i.e. either in the pulmonary artery or aorta). In this context, we believe that a similar modification can be made to accommodate tri-leaflet valves in other pulsatile simulators as well. Specifically in our case, to house a tri-leaflet valve structure in the aortic valve location, an assembly primarily of acrylic plastic (Plexiglass) casing that contained a rapid prototype valve holder and sutured tri-leaflet valve (Figures 1b and 1c) could be easily integrated and removed from the primary ViVitro system. Hydrodynamic testing was subsequently conducted similar to other studies performed by Baldwin et al.26 and Wang et al.25 Instantaneous flow rate was measured using an electromagnetic flowmeter system (Figure 3). Real-time measurement of pressure was recorded at the ventricular and conduit location using microtip transducers at a set heart rate of 70 beats/min (Figures 4 and 5). The testing fluid was a blood-analog liquid, comprising deionized water to glycerin in a 65% to 35% ratio and 9 g/L of NaCl, mimicking blood viscosity (~3.3 cP).
We initially tested a mechanical bi-leaflet valve and the obtained mean pressure wave forms were compared to literature values19. Some ventricular pressure variability was observed possibly owing to different pump mechanisms in place to drive fluid flow as well as geometry and specific settings of different pulse duplicator systems such as size of the ventricular chamber, specific valve mimicking the mitral valve location, heart rate chosen, physiological flow waveform selected, etc. On the other hand, the aortic waveforms were found to be very similar and system-independent. This exercise was repeated for native porcine valves and again, larger variability in ventricular pressure was observed when comparing our results to the literature18. However, it is important to note that within our system, instantaneous flow rates as well as both ventricular and aortic pressures were similar regardless of the valve that was tested, i.e. polymer and native with assembly or bi-leaflet without assembly. This exercise is important to perform because one needs to ensure that modifications to the duplicator system with an assembly do not considerably alter local flow and/or pressure conditions. Secondly, these results indicate that as a means of system validation, at minimum, comparable aortic pressures need to be derived across pulse duplicator platforms or the valve being tested. The interpretation of the hydrodynamic variables themselves is a matter of individual polymer valve design specifics. Standards such as ISO (International Organization for Standardization) 5840 used in the evaluation of cardiac valve prostheses can serve as a guide to assess various parameters associated with the polymer valve geometry, manufacturing and material properties. These parameters can be further optimized and hydrodynamic testing subsequently revisited to ensure that the standards needed for FDA submission are met.
For example, in our polymer valves, comparable energy losses and lower regurgitant volumes versus native and bi-leaflet valves suggested acceptable workloads on the left ventricle21 and efficient valve closure (Table 1). However, the closing dynamics resulted in a relatively higher polymer valve maximum TVP gradient (versus bi-leaflet valves), which in our case, warrants further mechanical evaluation of silicone material being used to fabricate the valves to ensure that the higher stress does not cause leaflet rupture, and that a sufficient factor of safety can be put in place. In conclusion, we have demonstrated that an assembly consisting of a housing unit, glass tube and a valve holder can be fabricated to accommodate tri-leaflet structures such as polymer valves which can be sutured in-position. Comparative flow and pressure waveforms across native, prosthetic and polymer valves that are being developed need to be obtained. Second, the pressure waveforms need to be validated with literature values. A limitation of our approach is that ventricular waveforms are pulse duplicator system specific and are likely to show differences; however aortic pressure waveforms should be comparable across platforms or valve being tested if sufficient valve functionality exists. A future direction of this work is to further optimize the polymer valve material, manufacturing process and geometry. Hydrodynamics tests will subsequently be repeated under identical conditions so as to determine if functional improvements are quantitatively observed by comparing the current and previous hydrodynamic metrics computed.
The authors have nothing to disclose.
A seed grant from the University of Florida – College of Medicine is gratefully acknowledged. Graduate studies (Manuel Salinas) were supported through a minority opportunities in biomedical research programs – research initiative for scientific enhancement (MBRS-RISE) fellowship: NIH/NIGMS R25 GM061347. Financial support from the Wallace H. Coulter Foundation through Florida International University’s, Biomedical Engineering Department is also gratefully acknowledged. Finally, the authors thank the following students for their assistance during various stages of the experimental process: Kamau Pier, Malachi Suttle, Kendall Armstrong and Abraham Alfonso.
Name of Reagent/Material | Company | Catalog Number | Comments |
Pump | ViVitro Labs | http://vivitrolabs.com/products/superpump/ | |
Flow Meter and Probe | Carolina Medical | Model 501D | http://www.carolinamedicalelectronics.com/documents/FM501.pdf |
Pressure Transducer | ViVitro Labs | HCM018 | |
ViVitro Pressure Measuring Assembly | ViVitro Labs | 6186 | |
Valve holder | WB Engineering | Designed by Florida International University. Manufactured by WB Engineering | |
Pulse Duplicator | ViVitro Labs | PD2010 | http://vivitrolabs.com/wp-content/uploads/Pulse-Duplicator-Accessories1.pdf |
Pulse Duplicator Data Acquisition and Control System, including ViViTest Software | ViVitro Labs | PDA2010 | http://vivitrolabs.com/products/software-daq |
Porcine Hearts and Native Aortic Valves | Mary's Ranch Inc | ||
Bi-leaflet Mechanical Valves | Saint Jude Medical | http://www.sjm.com/ | |
High Vacuum Grease | Dow Corning Corporation | http://www1.dowcorning.com/DataFiles/090007b281afed0e.pdf | |
Glycerin | McMaster-Carr | 3190K293 | 99% Natural 5 gal |
Phosphate Buffered Saline (PBS) | Fisher Scientific | MT21031CV | 100 ml/heart |
Antimycotic/Antibiotic Solution | Fisher Scientific | SV3007901 | 1 ml in 100 ml of PBS/heart; 20 ml for ViVitro System |
NaCl | Sigma-Aldrich | S3014-500G | 9 g/L of deionized water |
Deionized Water | EMD Millipore Chemicals | Millipore Deionized Purification System. 1.3 L for ViVitro System, 200 ml for heart valve dissection process |