- Email: [email protected]
Optimization of pulsatile flow for mechanical circulatory support
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EDITORIAL COMMENTARY
Optimization of pulsatile flow for mechanical circulatory support Michael J. Domanski, MD,a and Don P. Giddens, PhDb From the aZena and Michael Weiner Cardiovascular Institute, Mount Sinai School of Medicine, New York, New York; and the bWallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology, Atlanta, Georgia.
Current “continuous-flow” left ventricular assist devices (LVADs) prolong the survival of appropriately selected patients and improve their quality of life.1,2 However, these devices have been associated with a number of complications, possibly due to the non-physiologic pressure–flow relationships that they generate. These include arteriovenous malformations, gastrointestinal bleeding and aortic insufficiency.3,4 A reasonable hypothesis is that flow more closely mimicking that seen in the normal cardiovascular system could eliminate these complications. Designing a device to do this requires being able to quantify the biologically relevant parameters that characterize normal physiologic flow. In this issue of the journal, Soucy et al5 address the metrics of pulsatility. We join that discussion by starting from the initial principles. A flow field may be “steady,” meaning that it is invariant with time. Therefore, “unsteady” flow is flow that varies over time. A special case of unsteady flow is “pulsatile” flow, in which the flow field repeats itself rhythmically over time. Continuous-flow LVADs are basically turbines in a conduit, so that, in the absence of any cardiac contraction, they would produce a steady flow. Because most patients have some residual contractile function, a component of unsteady (although not necessarily pulsatile according to the definition presented) flow is superposed on the underlying and generally numerically dominant steady flow. Generating a pulsatile flow that mimics physiologic pulsatility requires a flow field that duplicates as closely as possible the characteristics of normal physiologic flow actually transduced (measured) at the cellular level. So, the question becomes how does the flow communicate with the cells of the vasculature? Or, what do the cells actually sense or measure? Reprint requests: Michael J. Domanski, MD, Cardiovascular Institute, Mount Sinai School of Medicine, One Gustav L. Levy Place, New York, NY 20514. Tel.: 646-537-8560. E-mail address: [email protected]
Vascular endothelial cells sense the fluid dynamic wall shear stress (WSS) exquisitely, and a number of studies have shown that they are sensitive to both mean (timeaveraged) WSS and also to oscillatory WSS.6 Indeed, much study has been devoted to the response of endothelial cells of various species to different steady- and pulsatile-flow conditions. As an example, when there is low mean WSS (on the order of o5 to 10 dynes/cm2), endothelial cells exhibit numerous pro-atherogenic and inflammatory responses.7 If an oscillatory WSS is superposed on the low mean WSS, the endothelial response can be somewhat different, but is still pro-atherogenic. In contrast, a “physiologic” WSS, typically considered to be 15 to 25 dynes/cm2, leads to an athero-protective response. If, on other hand, WSS is significantly greater than physiologic, then endothelial cells respond by causing the arterial wall to remodel and develop a larger lumen.8 This behavior makes teleological sense. For instance, when a child grows, there is an increased demand for blood flow to tissues, and there must be a mechanism to signal the appropriate cellular responses so that arteries become larger to accommodate the increased blood flow. Sensing WSS is an exquisitely sensitive way of regulating cellular behavior, because, as a rough approximation, WSS Q/D3, where Q is mean flow and D is vessel diameter. Thus, relatively small changes in D can accommodate increased Q and keep WSS in the physiologic range. The inverse is also true. A WSS lower than physiologic seems to signal the cells to remodel the artery wall to become smaller (e.g., intimal hyperplasia), presumably for maintaining a physiologic WSS for the endothelial cells (ECs). If the remodeling takes place in the presence of substances that are associated with risk for atherosclerosis, such as low-density lipoprotein (LDL) and smoking byproducts, then the remodeling can become atherosclerotic. In summary, the physiologic range of WSS for humans seems to be about 15 to 25 dynes/cm2. If mean WSS is lower, constrictive remodeling occurs, whereas, if WSS is higher, expansive remodeling takes
1053-2498/$ - see front matter r 2013 International Society for Heart and Lung Transplantation. All rights reserved. http://dx.doi.org/10.1016/j.healun.2013.03.013
578
The Journal of Heart and Lung Transplantation, Vol 32, No 6, June 2013
place (Glagov remodeling).9 In particular, low mean WSS elicits pro-atherogenic EC responses. Another consideration is related to avoiding regions of flow stasis, which can contribute to enhancing thrombosis in the short term and intimal hyperplasia (through the WSS mechanism described earlier) in the long term. In the vasculature, as well as in the geometric design of mechanical assist devices, there are unavoidably regions that are subject to stasis under steady-flow conditions because of local geometry, such as branches and bifurcations in vessels and abrupt changes in conduit area in devices. Out-of-plane curvature, which induces secondary or swirling flows, can reduce zones of stasis, even under steady-flow conditions, as can periodic alterations in continuous-flow settings, but the addition of pulsatility is more effective.10 There are also effects of pulsatility on smooth muscle cells. The pulsatile pressure–based stresses, such as hoop stress, within the artery wall are 5 orders of magnitude greater (1 106 dynes/cm2) than the fluid dynamic WSS acting on ECs, and smooth muscle cells respond differently to different pressure–time relationships.11,12 For instance, greater-than-normal pressure pulsatility can lead to vessel hypertrophy.13 Cyclic stretching of smooth muscle cells in culture results in a 2- to 5-fold increase in protein and collagen synthesis.14 A full discussion of cellular transduction of pressure, shear stress and the time and directional changes in these quantities is not possible in this short editorial commentary. The message is that the specifics of how a flow field varies over time are critically important in attempting to mimic physiologic flow. As pointed out by Soucy et al, Fourier analysis is a convenient approach to mathematical description of the flow–time relationship, which uses the sum of a series of sine and/or cosine functions to represent instantaneous values of pressure, flow, WSS and, for that matter, any other relevant hemodynamic quantity. The “average flow” component is the zeroth order (first) term in the Fourier series, and the subsequent terms (harmonics) are calculated at the fundamental heart rate and multiples thereof, effectively capturing the shape of the waveform in question. Soucy and colleagues used 10 harmonics, which should be more than adequate, because the fraction of the flow energy represented declines with each increasingly higher harmonic. To summarize, logic and biologic evidence suggest the importance of providing pressure and flow waveforms that result in WSS and wall stress–time relationships that are in the physiologic range. The metrics suggested by Soucy and colleagues (energy equivalent pressure, surplus hemodynamic energy) do
provide measures that are sensitive to pulsatile flow. However, these are summary statistics that could result from a variety of waveforms (and thus different timevarying relationships), including ones very different from those encountered physiologically. For this reason, these metrics may add value beyond other simple metrics that have been used previously, but we believe a more effective approach is to mimic parameters actually transduced by the cells communicating with the blood flow.
References 1. Mehra M, Domanski M. Should left ventricular assist device should be standard of care for patients with refractory heart failure who are not transplantation candidates? Circulation 2012;126:3081-7. 2. Rose EA, Gelijns AC, Moskowitz AJ, et al. Randomized Evaluation of Mechanical Assistance for the Treatment of Congestive Heart Failure (REMATCH) study group. Long-term use of a left ventricular assist device for end-stage heart failure. N Engl J Med 2001;345:1435-43. 3. Demirozu Z, Radovancevic R, Hochman L, et al. Arteriovenous malformation and gastrointestinal bleeding in patients with the HeartMate II left ventricular assist device. J Heart Lung Transplant 2011;30:849-53. 4. Martina J, de Jonge N, Sukkel E, et al. Left ventricular assist devicerelated systolic aortic regurgitation. Circulation 2011;124:487-8. 5. Soucy KG, Koenig SC, Giridharan GA. Defining pulsatility during continuous-flow ventricular assist device support. J Heart Lung Transplant 2013;32:581-7. 6. Rezvan A, Chih-Wen N, Alberts-Grill N, et al. Animal, in vitro, and ex vivo models of flow-dependent atherosclerosis: role of oxidative stress. Antioxid Redox Signal 2011;15:1433-48. 7. Uzarski J, Scott E, McFetridge PS. Adaptation of endothelial cells to physiologically modeled, variable shear stress. PLoS One 2013;8: e57004. 8. Zairns C, Zatina M, Giddens D, et al. Shear stress regulation of artery lumen diameter in experimental atherogenesis. J Vasc Surg 1987; 5:413-20. 9. Glagov S, Weisenberg E, Zarins C, et al. Compensatory enlargement of human atherosclerotic coronary arteries. N Engl J Med 1987; 316:1371-5. 10. Giddens E, Giddens D, White S, et al. Exercise flow conditions eliminate stasis at vascular graft anastomoses. In: Schneck D, Lewis C, editors. Biomechanics 3. Proceedings of the Third Mid-Atlantic Conference in Biofluid Mechanics. New York: New York University Press 1990;255-67. 11. Haga J, Li Y, Chien S. Molecular basis of the effects of mechanical stretch on vascular smooth muscle cells (review). J Biomech 2007;40:947-60. 12. Laurent S, Tropeano A, Lillo-Lelouet A, et al. Local pulse pressure is a major determinant of large artery remodeling. Clin Exp Pharmacol Physiol 2001;28:1011-4. 13. Eberth J, Gresham VC, Reddy AK, et al. Importance of pulsatility in hypertensive carotid artery growth and remodeling. J Hypertens 2009;27:2010-21. 14. Sottiurai V, Kollros P, Glagov S, et al. Morphologic alteration of cultured arterial smooth muscle cells by cyclic stretching. J Surg Res 1983;35:490-7.
EDITORIAL COMMENTARY
Optimization of pulsatile flow for mechanical circulatory support Michael J. Domanski, MD,a and Don P. Giddens, PhDb From the aZena and Michael Weiner Cardiovascular Institute, Mount Sinai School of Medicine, New York, New York; and the bWallace H. Coulter Department of Biomedical Engineering, Georgia Institute of Technology, Atlanta, Georgia.
Current “continuous-flow” left ventricular assist devices (LVADs) prolong the survival of appropriately selected patients and improve their quality of life.1,2 However, these devices have been associated with a number of complications, possibly due to the non-physiologic pressure–flow relationships that they generate. These include arteriovenous malformations, gastrointestinal bleeding and aortic insufficiency.3,4 A reasonable hypothesis is that flow more closely mimicking that seen in the normal cardiovascular system could eliminate these complications. Designing a device to do this requires being able to quantify the biologically relevant parameters that characterize normal physiologic flow. In this issue of the journal, Soucy et al5 address the metrics of pulsatility. We join that discussion by starting from the initial principles. A flow field may be “steady,” meaning that it is invariant with time. Therefore, “unsteady” flow is flow that varies over time. A special case of unsteady flow is “pulsatile” flow, in which the flow field repeats itself rhythmically over time. Continuous-flow LVADs are basically turbines in a conduit, so that, in the absence of any cardiac contraction, they would produce a steady flow. Because most patients have some residual contractile function, a component of unsteady (although not necessarily pulsatile according to the definition presented) flow is superposed on the underlying and generally numerically dominant steady flow. Generating a pulsatile flow that mimics physiologic pulsatility requires a flow field that duplicates as closely as possible the characteristics of normal physiologic flow actually transduced (measured) at the cellular level. So, the question becomes how does the flow communicate with the cells of the vasculature? Or, what do the cells actually sense or measure? Reprint requests: Michael J. Domanski, MD, Cardiovascular Institute, Mount Sinai School of Medicine, One Gustav L. Levy Place, New York, NY 20514. Tel.: 646-537-8560. E-mail address: [email protected]
Vascular endothelial cells sense the fluid dynamic wall shear stress (WSS) exquisitely, and a number of studies have shown that they are sensitive to both mean (timeaveraged) WSS and also to oscillatory WSS.6 Indeed, much study has been devoted to the response of endothelial cells of various species to different steady- and pulsatile-flow conditions. As an example, when there is low mean WSS (on the order of o5 to 10 dynes/cm2), endothelial cells exhibit numerous pro-atherogenic and inflammatory responses.7 If an oscillatory WSS is superposed on the low mean WSS, the endothelial response can be somewhat different, but is still pro-atherogenic. In contrast, a “physiologic” WSS, typically considered to be 15 to 25 dynes/cm2, leads to an athero-protective response. If, on other hand, WSS is significantly greater than physiologic, then endothelial cells respond by causing the arterial wall to remodel and develop a larger lumen.8 This behavior makes teleological sense. For instance, when a child grows, there is an increased demand for blood flow to tissues, and there must be a mechanism to signal the appropriate cellular responses so that arteries become larger to accommodate the increased blood flow. Sensing WSS is an exquisitely sensitive way of regulating cellular behavior, because, as a rough approximation, WSS Q/D3, where Q is mean flow and D is vessel diameter. Thus, relatively small changes in D can accommodate increased Q and keep WSS in the physiologic range. The inverse is also true. A WSS lower than physiologic seems to signal the cells to remodel the artery wall to become smaller (e.g., intimal hyperplasia), presumably for maintaining a physiologic WSS for the endothelial cells (ECs). If the remodeling takes place in the presence of substances that are associated with risk for atherosclerosis, such as low-density lipoprotein (LDL) and smoking byproducts, then the remodeling can become atherosclerotic. In summary, the physiologic range of WSS for humans seems to be about 15 to 25 dynes/cm2. If mean WSS is lower, constrictive remodeling occurs, whereas, if WSS is higher, expansive remodeling takes
1053-2498/$ - see front matter r 2013 International Society for Heart and Lung Transplantation. All rights reserved. http://dx.doi.org/10.1016/j.healun.2013.03.013
578
The Journal of Heart and Lung Transplantation, Vol 32, No 6, June 2013
place (Glagov remodeling).9 In particular, low mean WSS elicits pro-atherogenic EC responses. Another consideration is related to avoiding regions of flow stasis, which can contribute to enhancing thrombosis in the short term and intimal hyperplasia (through the WSS mechanism described earlier) in the long term. In the vasculature, as well as in the geometric design of mechanical assist devices, there are unavoidably regions that are subject to stasis under steady-flow conditions because of local geometry, such as branches and bifurcations in vessels and abrupt changes in conduit area in devices. Out-of-plane curvature, which induces secondary or swirling flows, can reduce zones of stasis, even under steady-flow conditions, as can periodic alterations in continuous-flow settings, but the addition of pulsatility is more effective.10 There are also effects of pulsatility on smooth muscle cells. The pulsatile pressure–based stresses, such as hoop stress, within the artery wall are 5 orders of magnitude greater (1 106 dynes/cm2) than the fluid dynamic WSS acting on ECs, and smooth muscle cells respond differently to different pressure–time relationships.11,12 For instance, greater-than-normal pressure pulsatility can lead to vessel hypertrophy.13 Cyclic stretching of smooth muscle cells in culture results in a 2- to 5-fold increase in protein and collagen synthesis.14 A full discussion of cellular transduction of pressure, shear stress and the time and directional changes in these quantities is not possible in this short editorial commentary. The message is that the specifics of how a flow field varies over time are critically important in attempting to mimic physiologic flow. As pointed out by Soucy et al, Fourier analysis is a convenient approach to mathematical description of the flow–time relationship, which uses the sum of a series of sine and/or cosine functions to represent instantaneous values of pressure, flow, WSS and, for that matter, any other relevant hemodynamic quantity. The “average flow” component is the zeroth order (first) term in the Fourier series, and the subsequent terms (harmonics) are calculated at the fundamental heart rate and multiples thereof, effectively capturing the shape of the waveform in question. Soucy and colleagues used 10 harmonics, which should be more than adequate, because the fraction of the flow energy represented declines with each increasingly higher harmonic. To summarize, logic and biologic evidence suggest the importance of providing pressure and flow waveforms that result in WSS and wall stress–time relationships that are in the physiologic range. The metrics suggested by Soucy and colleagues (energy equivalent pressure, surplus hemodynamic energy) do
provide measures that are sensitive to pulsatile flow. However, these are summary statistics that could result from a variety of waveforms (and thus different timevarying relationships), including ones very different from those encountered physiologically. For this reason, these metrics may add value beyond other simple metrics that have been used previously, but we believe a more effective approach is to mimic parameters actually transduced by the cells communicating with the blood flow.
References 1. Mehra M, Domanski M. Should left ventricular assist device should be standard of care for patients with refractory heart failure who are not transplantation candidates? Circulation 2012;126:3081-7. 2. Rose EA, Gelijns AC, Moskowitz AJ, et al. Randomized Evaluation of Mechanical Assistance for the Treatment of Congestive Heart Failure (REMATCH) study group. Long-term use of a left ventricular assist device for end-stage heart failure. N Engl J Med 2001;345:1435-43. 3. Demirozu Z, Radovancevic R, Hochman L, et al. Arteriovenous malformation and gastrointestinal bleeding in patients with the HeartMate II left ventricular assist device. J Heart Lung Transplant 2011;30:849-53. 4. Martina J, de Jonge N, Sukkel E, et al. Left ventricular assist devicerelated systolic aortic regurgitation. Circulation 2011;124:487-8. 5. Soucy KG, Koenig SC, Giridharan GA. Defining pulsatility during continuous-flow ventricular assist device support. J Heart Lung Transplant 2013;32:581-7. 6. Rezvan A, Chih-Wen N, Alberts-Grill N, et al. Animal, in vitro, and ex vivo models of flow-dependent atherosclerosis: role of oxidative stress. Antioxid Redox Signal 2011;15:1433-48. 7. Uzarski J, Scott E, McFetridge PS. Adaptation of endothelial cells to physiologically modeled, variable shear stress. PLoS One 2013;8: e57004. 8. Zairns C, Zatina M, Giddens D, et al. Shear stress regulation of artery lumen diameter in experimental atherogenesis. J Vasc Surg 1987; 5:413-20. 9. Glagov S, Weisenberg E, Zarins C, et al. Compensatory enlargement of human atherosclerotic coronary arteries. N Engl J Med 1987; 316:1371-5. 10. Giddens E, Giddens D, White S, et al. Exercise flow conditions eliminate stasis at vascular graft anastomoses. In: Schneck D, Lewis C, editors. Biomechanics 3. Proceedings of the Third Mid-Atlantic Conference in Biofluid Mechanics. New York: New York University Press 1990;255-67. 11. Haga J, Li Y, Chien S. Molecular basis of the effects of mechanical stretch on vascular smooth muscle cells (review). J Biomech 2007;40:947-60. 12. Laurent S, Tropeano A, Lillo-Lelouet A, et al. Local pulse pressure is a major determinant of large artery remodeling. Clin Exp Pharmacol Physiol 2001;28:1011-4. 13. Eberth J, Gresham VC, Reddy AK, et al. Importance of pulsatility in hypertensive carotid artery growth and remodeling. J Hypertens 2009;27:2010-21. 14. Sottiurai V, Kollros P, Glagov S, et al. Morphologic alteration of cultured arterial smooth muscle cells by cyclic stretching. J Surg Res 1983;35:490-7.