Similar to the population with the lowest organ availability, 40% of patients younger than 1 year will die while awaiting a heart transplant.1 In many cases, these pediatric patients would benefit from circulatory support to bridge the time to transplantation; however, development of such systems has not progressed as fast as that for the much larger adult population.2 To rectify this, the National Heart, Lung and Blood Institute developed and initiated the Pediatric Circulatory Support System Program that aims to develop support options for patients with body weight ranging from 2 to 25 kg.3 As part of this program, Penn State is currently developing 12 and 25 cc pediatric ventricular assist devices (PVADs) based on the successful pneumatically driven Pierce-Donachy adult device, currently used clinically as the Thoratec.4
Although the use of circulatory support in patients increases the chance of pretransplantation survival, there are concerns in using such systems for extended periods of time. For example, extracorporeal membrane oxygenation, currently the most used form of circulatory support for pediatric patients, can cause blood damage, infection, and bleeding when used for >72 hours.5 For ventricular assist devices (VADs), one of the major obstacles to long-term use is blood damage and a subsequent increase in thrombus formation and detachment.6 This issue can be exacerbated by poor fluid dynamic washing inside the device. Previous studies of pulsatile VADs, similar to the Penn State PVAD, have found that desirable flow conditions for the reduction of thrombus deposition include a strong diastolic inlet jet followed by a continuous rotational flow without areas of stagnation.7 Furthermore, work by Hubbell and McIntire8 found that shear rates >500 s−1 on polyurethane materials, such as that used for the blood sac of the PVAD, are sufficient to prevent thrombus formation.
The fluid mechanics of the Penn State PVAD have been investigated indepth using in vitro flow visualization techniques.9–14 These studies have indicated that the reduction in volume necessary to accommodate pediatric patients can lead to undesirable changes in the flow. These include an increase in the three-dimensionality of the flow and areas of separation and regurgitation, both of which are strongly coupled to the flow through the valves. This brings into question how changes in the valve orientation may affect the PVAD flow field.
Previous studies of a pulsatile 50 cc adult VAD found that changing the orientation of the inlet valve by 30° led to an improved inlet jet flow, including an increase both in the jet penetration and in the magnitude and duration of the wall shear along the device walls.15 A flow visualization study of a 75 cc device by Akagawa et al.16 showed that a change in orientation of the inlet valve by 45° led to an increase in the inlet flow velocity and helped prevent flow stagnation. Although both these studies focused solely on the inlet valve orientation, the much smaller volume of the PVAD makes the flow field more sensitive to valve-related effects, so that when the orientations of both the inlet and outlet valves are varied independently, or together, may signify different flow fields.
Materials and Methods
To perform in vitro flow visualization, an optically transparent model of the PVAD was constructed as shown in Figure 1. The acrylic model was contoured to match the curvature of the PVAD blood sac. The diaphragm encompassed one side of the device, and the inlet and outlet ports were angled 20° away from this side of the device. The model was fitted with 17 mm Björk-Shiley Monostrut (BSM) mechanical heart valves. The initial orientation (0°) aligned the opening of the major orifice of the BSM valve along the outer wall of the device on both the inlet and outlet sides. The valves were then rotated toward either the diaphragm or fluid side of the device. For this study, orientation toward the fluid side of the device was defined as positive, whereas orientation toward the diaphragm side was defined as negative. Figure 2 presents examples of these orientations, and Table 1 lists all the orientations evaluated.
To simulate physiologic conditions, the PVAD was placed in a mock circulatory loop originally developed by Rosenberg et al.17 to study VADs and artificial hearts. We made slight modifications to the loop for this pediatric application, including an increase in the resistance and a reduction in the overall volume. A viscoelastic fluid developed to match the properties of 40% hematocrit pediatric blood was used.18 The fluid contained Xanthan gum (0.03%), glycerin (16%), sodium iodide (50%), and water (35%) by weight. The fluid had a refractive index of 1.488 to match that of the acrylic model and, to permit flow visualization, was seeded with 10-μm hollow glass beads.
A pneumatic driver powered the PVAD at a rate of 75 beats per minute with a systolic duration of 340 msec. A flow rate of 1.4 L/min was maintained, and the flow waveforms at the inlet and outlet of the device were measured using ultrasonic flow probes (Transonic Systems, Inc., Millis, MA). Atrial (inlet) and aortic (outlet) pressures were measured using pressure transducers (Maxxim Medical, Athens, TX). All flow and pressure signals were converted and displayed using a WaveBook acquisition system and WaveView software (IOtech, Inc., Cleveland, OH). An example of the flow and pressure waveforms maintained is shown in Figure 3.
Particle image velocimetry (PIV) was performed in three planes as shown in Figure 4A. These planes were parallel to the device diaphragm at 7, 8.2, and 11 mm from the edge of the port. The 7- and 8.2-mm planes allowed for observation into the ports while still providing flow information in the body of the device while the 11-mm plane encompassed the entire body of the device. Frequent diaphragm interference prevented complete study of the flow field further into the device.
A Gemini PIV 15 system (New Wave Research, Inc., Fremont, CA) was used and included dual Nd:YAG lasers and associated optics to form a 200-μm light sheet. Images were captured using a 1 megapixel charge-coupled device camera (TSI, Inc., Shoreview, MN) with a Micro-Nikkor 60-mm f/2.8D lens (Nikon Corporation, Tokyo, Japan). Two hundred images were acquired at time steps 50 msec apart throughout the cardiac cycle resulting in 15 data sets for one cardiac cycle. These images were then processed using MATLAB 7.1 (The MathWorks, Inc., Natick, MA) and Insight 3G software (TSI, Inc., Shoreview, MN). Previous studies by Cooper et al.12,13 and Roszelle et al.11,14 fully describe the details of the PIV used in this study. In addition, an error analysis is given by Cooper et al.12 The process for obtaining wall shear rates required finding the fluid boundary or “walls” of the device (Figure 4B) and used a centroid shifting technique to obtain the near wall velocities. This process was first described by Hochareon et al.19
Inlet Valve Orientation
Because the inlet valve angle was changed, the rotational flow pattern observed in previous studies10–14 was generally maintained. The major changes to the flow field occurred at the diastolic inlet jet. Because the major orifice of the valve was turned toward the fluid side of the PVAD, the jet spread parallel to the diaphragm, and there was an increased flow through the minor orifice (Figure 5). This spreading resulted in an increase in wall shear rates above the 500 s−1 threshold over a larger region across surface 1 (Figure 4B). Figure 6 highlights this increase in wall shear, showing for the 7-mm plane; the percentage of surface 1 with shear >500 s−1 increases from 40% to 65% to 100% as the valve angle increases from 0° to +15° to +30°, respectively. However, at an angle of +45°, the shear was lower and only occurred over 35% of surface 1 (Figure 6). As highlighted by a red oval in Figure 5, at an inlet valve angle of 0°, there was an area of flow discontinuity between the inlet jet and the beginning of the rotational flow at the bottom of the device, which is characteristic of the three-dimensional flow seen in the PVAD.14 As the valve angle was turned toward the fluid side, this flow discontinuity decreases, leading to a more cohesive rotational flow. However, as highlighted in Figure 5, at +45°, this discontinuity resurfaces. Previous studies have shown that an increase in three-dimensionality, or flow normal to the diaphragm, affects the overall development of the inlet jet by blocking this penetration to the bottom of the device.4,9,14 The more cohesive flow found with the angle increase indicated better penetration, which was also confirmed with increased wall shear rates along the outer inlet wall as seen in Figure 6. With an increase in the positive orientation of the valve, there was also a greater influence of the minor orifice jet. Although the minor orifice jet did not have a significant effect on the rotational flow in angles up to +30°, at +45°, it became dominant, and the combination of the wider major orifice inlet jet and a larger minor orifice jet led to a less organized rotational flow. In comparison with the positive angles, when angled at −15° (the major orifice of the valve shifted toward the diaphragm side), there was a large separation in the flow about the inlet valve. Along with this separation, there was greater three-dimensionality in the flow, which, for the +45° orientation, was highlighted by a large area of discontinuity between the diastolic jet and the start of the rotational flow field (Figure 5).
Outlet Valve Orientation
Because the initial observations of the inlet valve changes showed the best results between 0° and +30°, this range was used for the outlet valve orientation. When the inlet valve was at an orientation of 0° and the outlet valve was rotated between 0° and +30°, the diastolic flow patterns were nearly identical, except for a slight change in magnitude of the inlet jet. Some differences were observed between the orientations during systole. An increase in valve angle corresponded to an increase in the wall shear rate along surface 4 (Figure 4B) as identified in Figure 7. Both the shear rate magnitude and the area for which the shear rates were >500 s−1 increased. A unique characteristic seen previously during late systole with the BSM valve was an area of blockage upstream of the outlet valve. As the angle of the valve was increased, this area of blockage remained and showed a slight increase in size as shown in Figure 8.
Combined Valve Orientation
When both the inlet and the outlet valve angles were increased simultaneously, the flow exhibited the pattern that occurred with rotating the valves individually. Because our major concern is the wall shear, we present the comparison in the form of wall shear rate maps. Figure 9 compares the wall shear rates along the entire outer surface of the PVAD at the 11-mm plane for all four combined valve orientations (Figure 4B highlights these surface locations). Although a large percentage of the walls for each valve combination maintained adequate shear (>500 s−1 during some point of the cycle), the (+15°, +15°) orientation had the most desirable wall shear rates. At surface 1, the closest to the inlet valve, every combination except for (+30°, +15°) had shear well above the threshold value for at least 300 msec. For surface 2, however, the (+15°, +15°) combination had shear rates above the threshold value both for the largest portion of the wall and for the longest part of cycle. This is more apparent when compared with the (+30°, +15°) and the (+30°, +30°) combinations. Each combination showed adequate wall shear rates at surface 3; however, the (+15°, +15°) and (+15°, +30°) combinations had a longer duration of adequate shear rates and a more uniform pattern. At surface 4, we again observed adequate shear in each combination during systole except for the (+30°, +15°) orientation. All four combinations also show a low shear region at the edge of surface 4, which could be due to flow regurgitation.
In general, changes in valve orientation between 0° and +30° did not result in dramatically different flow patterns inside of the body of the PVAD. In that sense, these results are similar to those in a study by Avarahmi, which compared different valve types in numerical models of a pulsatile adult VAD. In their study, the flow patterns inside the VAD showed little change with valve type; however, the shear stresses and wash out properties, governed more by the flow at the wall, were affected by these changes.20 In our study, we found that valve orientations between 0° and +30° all produced a strong inlet diastolic jet that transitions to a rotational flow pattern in the body of the device through diastole and into systole. With the changing orientations, however, important changes occurred in the wall shear rates, with some combinations resulting in shear rates above the 500 s−1 threshold for both longer durations and greater portions of the sac wall. Because of the importance of reducing thrombus formation, these changes in shear may be significant.
Positive shifts in the inlet valve orientation, which move the flow away from the diaphragm, led to an increase in the wall shear rates along the inlet side of the device. This observation is consistent with data from Kreider et al., which found that increasing the positive angle of the valve led to higher near wall velocities on the inlet side of a 50 cc adult VAD. The study by Kreider et al.15 and a similar study by Akagawa et al.16 both found the ideal inlet valve position to be +45°. In contrast, our data suggest that a valve angle of +45° is too large for best performance. Akagawa et al. also found that the rotation of the valve toward the diaphragm (−45°) had the same positive effects as a rotation toward the fluid side of the device. Again, we did not observe this in the PVAD, because the rotation of the inlet valve to −15° led to a less cohesive rotation and a reduction in wall shear rates. Both differences are perhaps due to the smaller volume and angled ports of the PVAD. Because previous studies have shown that the reduced volume has led to an increased sensitivity of the flow field,9–14 the small changes in valve orientation have a greater impact, making the 45° angle too large. On the whole, we found that shifting the inlet valve toward the fluid side of the device up to +30° led to better wall shear rates along the inlet and a tighter rotational flow field, all desirable qualities for pulsatile VAD flow fields.
Previous valve orientation studies of adult VADs did not consider changes to the outlet valve because of its limited influence on the flow field, because early work suggested that the diastolic phase was the most important influence on the flow of the pump. Herein, during systole, increasing the outlet valve angle led to an increase in the wall shear rates along the outer outlet wall and an increase in the size of the area of blockage upstream of the valve. The increased wall shear rates are a positive aspect of the increased valve angle and are similar to that observed for the inlet valve with an increased angle. Overall, our data indicate that increasing both valve angles up to +30° leads to improved wall shear rates, which may correspond to a reduction in thrombus formation. The increased size of the blockage at the outlet that correlates to the increase in valve angle may be of concern because it could lead to an increase in blood damage through the higher shear of a more confined outlet jet. However, the blockage occurs with all outlet angles (0°, +15°, and +30°), and the increased blockage does not appear large enough to outweigh the improved wall shear rates obtained at the outlet with this increased valve angle. It should also be noted that changes of orientation of the outlet valve affected the flow patterns during diastole. This observation emphasizes the sensitivity of the flow field and the communication between the inlet and outlet valve flows of the PVAD. With this in mind, our data show that the (+15°, +15°) combination resulted in the best wall shear rate duration and coverage, and we recommend these angles.
In our study, we observed the effect of changing the BSM valve orientations at both the inlet and outlet in the 12 cc Penn State PVAD. As in previous studies, a rotation of the inlet valve to the fluid side of the device resulted in higher inlet jet velocities and larger wall shear rates. For rotation of the outlet valve orientations, we show for the first time that a positive rotation (toward the fluid side) provides increased wall shear rate magnitudes and durations. These positive aspects of the individual valve rotation were maintained when the valves were shifted simultaneously. Overall, the data showed that all angles between 0° and +30° resulted in acceptable flow patterns and wall shear rates, with the (+15°, +15°) combination providing the best results. Angles greater than +30° and those orientated toward the diaphragm side (−15°) resulted in undesirable changes to the flow field physics. The study also highlighted the increased sensitivity of the flow field that has resulted from the reduction of volume and a resultant increase in the communication between the valves. It should also be noted that if this study had only observed the flow field patterns found through flow visualization studies, the differences between the valve orientations may have looked minor, and the changes in wall shear rates, which are an important part of device thrombogenicity, would have been overlooked. This highlights the importance of looking at local flow measurements, in this case, the wall shear rates, instead of only making whole field observations.
Supported by NIH NHLBI Grant HV 48191.
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