Axial flow artificial heart blood pumps: a brief review.
Historically the medical arts have found great application in the areas of communicable disease prevention and treatment as well as the treatment of injuries. In the second half of the last century much medical research shifted towards rectification of more sedentary diseases such as cardiovascular-related diseases, cancer. Heart transplantation represents the culmination of much of this work. Despite great advances in medications and surgical techniques heart transplantation still remains relatively uncommon due to the prohibitively high cost, limited donor heart availability and limited expertise to handle heart transplants [1, 2]. It has been estimated that in the US alone over 700,000 people with congestive heart failure (CHF) die each year, making it the single most common cause of mortality [1,3]. Even if the patient receives a heart transplant there are still high mortality rates due to organ rejection, operative trauma and complications . Long term immunosupression required with heart transplants also degrades the attractiveness of this option as well . One possible solution of reducing the mortality rates and increasing the success of heart transplants is to develop biomechanical devices such as the axial blood pump.
Total Artificial Heart (TAH)
The replacement of the ailing or the failing heart with a mechanical device has been a long term dream of cardiac surgeon for many years. Even with the option of heart transplants, the total artificial heart is seen as an option for sustaining the patient while waiting for a donor heart. Morris  credited the start of the U.S Artificial Heart program under President Johnson by the National Institute of Health. The first implementation of the total artificial heart on a human being was done on Barney Clark from Seattle whom was suffering from severe congestive heart failure.
After the implementation of the TAH, Clark lived for 112 days which marked the success of the functionality of total artificial heart on human beings. Morris  presented the review of artificial hearts on two different makes, namely CardioWest and AbioCor with the operating principles are deliberately explained. The CardioWest is an iteration of the original Jarvik-7 while AbioCor is built on thorasic units, both TAH are in principal perform the same function: to ensure the blood is pumped as to how a biological heart would function. [5,35,37] explained the distinctive advantages which has been noted, such as CardioWest TAH large valves and short outflow blood path are advantageous in decreasing statis and thrombosis, while AbioCor TAH energy converter which consists of pressurized hydraulic fluid which helps to create the systole effect which the beat rate and balance are set manually. The left and right motor are automatic in nature, and by enabling the rate to be controlled manually, the use can increase cardiac output and decrease right atrial pressure [5,36,38]. While the US Food and Drug Administration (FDA) approved the use of TAH in the 90s, such mechanical devices are still in clinical trial, such as the CardioWest which is currently undergoing clinical trials in the US. The advancements in TAH include the miniaturization of the device [6, 7,39,40], reducing the effect of thrombosis  and improving the valve material and design.
Left Ventricle Assist Devices
Given the size and complexity of earlier TAHs some researchers shifted their attention to smaller and simpler pumping mechanisms which could support the ailing heart of CHF patients rather than replace it. The left ventricle assist device (LVAD) is a pumping mechanism placed between the ascending aorta and the left ventricle of the heart, supplying pressurized blood to the left ventricle of the native heart. Most congestive heart failures are due to left ventricle failure in adults [6,41,44]. Beginning in the 1990s these second generation devices were intended for temporary support of an ailing heart until transplantation could be performed. It was soon realized, however, that with the assistance of an LVAD the native heart function often improved significantly. This reverse remodeling has resulted in longer term usage of LVADs, including usage as a destination therapy in patients ineligible for heart transplants . These second generation LVADs featured improved materials, and were smaller than the TAHs of the 1980s. A variety of mechanisms, both electrical and pneumatic were developed, including an increasing number of rotary pumps. Rotary pumps, either centrifugal or axial, have a spinning impeller pumping the blood continuously. Unless the pump speed is intentionally modulated these pumps do not produce any detectable pulse. The longevity of a rotary blood pump is mainly determined by the durability of its wearing mechanical parts such as bearings and seals [9,42,43].
Pulsatile versus Continuous Flow
Initially it was assumed the native heart is pulsatile in nature, a pulsed blood flow was the requirement for artificial blood pumping systems [4,42]. A pulsatile blood flow tends to clear away cell agglomerations at flow stagnation points, reducing clotting and the incidence of stroke and other associated problems. While some still argue that a pulsatile blood flow is required [1,41], there is mounting evidence that continuous flow is acceptable for durations of up to several years [3,4,10,11,12, 13,14,46]. Given the smaller size and simplicity of continuous flow blood pumps, the long-term overall performance has been shown to be superior to pulsatile blood pumps during invivo trials . Even with initially pulsatile blood flow, by the time blood reaches the capillaries the pulsatile amplitude has been reduced, and it is continuous .
Centrifugal Blood Pump
Centrifugal pumps have relatively flat, disk-shaped impellers with radial fins on the fluid side of the disk. Blood is drawn into the pump near the center of the impeller, and is centrifugally flung radially outwards by the spinning action of the impeller. The blood accumulates in a radial cavity where it is ejected from the pump tangential to the outer diameter of the impeller. Centrifugal pumps can attain physiological flow rates (5 liters/minute) and pressures (100 mHg) at impeller speeds of 2000 to 4000 revolutions per minute (RPM).
Axial Blood Pump
Axial pumps are more cylindrical-shaped and smaller, compared to centrifugal pumps. The impeller is a multifinned "screw" mechanism which rotates about the axis of the pump . Blood enters the pump and passes through an axially finned flow guide called a flow straightener. The straightener is stationary and its function is to direct the flow of blood towards the impeller without undue rotation of the entering blood. At the other end of the straightener, the blades may curve slightly in the direction of rotation of the impeller, reducing the shear stress at this point in the flow. The impeller is the only moving part of the pump. Blood entering the impeller is both rotated with the impeller, and pushed towards the outlet of the pump by the screw action of the impeller. A narrow gap separates the outer diameter of the impeller blades from the pump housing. Blood exiting the impeller has higher pressure than the inlet and is rotating in the sense of the impeller. It then passes to a diffuser section similar to the straightener at the inlet, but with more twist to the blades. The diffuser takes the rotating blood and smoothly bends the flow into the axial direction. Some of the rotational inertia of the blood is converted to axial momentum, raising the outlet pressure.
The pump housing is larger than the inlet/outlet veins as blood must pass through the radial space between the inner body of the pump components and the housing wall. Careful optimization of the inlet and outlet geometry is required to avoid areas of recirculation and stagnation, which might contribute to blood cell accumulation and thrombosis. Initial versions of rotary pumps relied on impellers supported by sealed ceramic bearings and shafts. This arrangement resulted in wear and contamination issues from the seals and bearings, as well as blood damage at the sliding interface. More recent "third generation" axial blood pumps have magnetically suspended impellers with imbedded permanent magnets. These newer designs are electrically configured as multiphase brushless direct current motors with impeller speeds ranging from 6000 to 12,000 RPM at physiological pressures and flow rates . The power required to operate the pump depends on the flow rate and the pressure across the pump, and is typically in the range of 5 to 18 W [17,35].
The control systems for these pumps are compact enough to allow implantation along with the pump. Transcutaneous power transmission enables implantable systems with no leads piercing the skin. A circular coil of wire is implanted just below the skin of the abdomen, allowing easy access. A similar shaped coil of wire is placed extracorperally in proximity to the implanted coil and energized with an alternating current. This forms an air-core transformer, which transmits power from the external coil to the embedded one, connected to electronic circuitry and power storage batteries as seen in table 1. The pump may thus be operated solely from the internal batteries for a period of up to a few hours , allowing complete freedom of motion. To charge the system, an external harness is worn which contains the charging coil connected either to the electrical power grid, or an external battery power unit. The ability to move about without connection to a fixed console allows greatly enhanced mobility and improvement in the quality of life of the patient [18,19]. Patients with such systems may be discharged from the hospital and treated via outpatient care, greatly reducing the overall cost .
The comparatively small size of the axial blood pump allows it to be used in special applications where larger centrifugal and pulsatile pumps are difficult to incorporate. While most TAHs and larger centrifugal pumps can only be implanted in larger male patients, axial blood pumps have been fabricated which are small enough for pediatric use . Another unique application for axial blood pumps is use as a catheter-based pump. These pumps are very small and spin at speeds of tens of thousands of RPM. This can be used for rapid insertion with minimal invasiveness when stabilizing a patient. Typically the device supplies enough blood flow to sustain life, but not enough for complete functioning . This technique is very useful for supporting a patient while deciding what to do long term.
Challenges in Axial Blood Pump Development
Many of the challenges facing axial blood pumps are common to the other artificial pumps as well. With any implanted device, there is trauma associated with the implantation and increased risk of infection in tissue surrounding the device. Researchers have made great strides in biocompatibility of materials and special surface textures which encourage natural cell growth reducing the risk of thrombosis . Minimizing the size of the implanted device reduces the trauma of implantation and risk of infection. Eliminating the percutanious power leads is a key factor in improving reliability and reducing the risk of infection and bleeding [10, 11]. Even with electromagnetic transcutaneous power transmission, implantation of a controller and battery system is still required. Obviously reducing the power required reduces the size and problems associated with a totally implantable system as well.
The reliability of any implanted device is another important aspect. With only one moving part, i.e. the impeller, the continuous flow blood pumps are simpler, and thus have the potential to be more reliable, compared to more complicated pulsatile devices requiring valves. The magnetically-levitated impeller of recent devices almost completely eliminates wear, insuring long term mechanical integrity of the pump mechanism . A recent 18-month trial on 281 patients with axial pump LVADs showed no failures of the mechanical pump . The main reasons for the growing interest in axial blood pump is their small size, power requirements and reliability [5(10)]. Being small size however, very high rotational speeds are required to deliver sufficient blood flow. While centrifugal pumps are generally operated at speeds below 4000 RPM, axial blood pumps are usually operated at 6000 to 12,000 RPM, with small catheter based systems spinning at over 30,000 RPM [2, 11, 23]. Hemolysis is the loss of haemoglobin and destruction of red blood cells caused by high shear stress . As pump speeds increase the velocity gradient in the fluid increases, thereby increasing the shear stress and hemolysis. The issue of hemolysis and blood cell destruction is generally worse with axial blood pumps than centrifugal blood pumps due to the higher speed of the axial pumps. Pulsatile pumps have been shown to have worse hemolysis than continuous flow pumps due to the existence of valves in the blood flow . A summary comparing some of the relevant parameters of pulsatile pumps, centrifugal and axial blood pumps is shown in table 1.
Another problem associated with artificial blood pumps is thrombosis. Thrombosis is the formation of a clot of blood cells which may be triggered by flow stagnation and abnormalities in the blood flow pattern or walls of the blood vessel. Thrombosis related incidents are one of the primary problems associated with axial blood pumps [1, 23].
Cardiovascular assist devices are prosthetic machines that are implanted into the body to assist the heart function. An artificial heart or VAD may be made out of animal tissue, plastic, ceramic and metal based on the required functional and mechanical properties. Titanium alloys are often used in VADs because of its biocompatibility and suitable mechanical properties (36,39,40). For blood interfacing surfaces, the titanium is machined to a specific degree of finish. The faces in contact with blood may receive a special coating of titanium microspheres that permanently bond to the surface. With this coating, blood cells stick to the surface. Alternatively a specific kind of polyurethane can be used as a surface coating to provide blood cell adherence. Polyester is often used in other parts of the system such as tubular grafts. The impeller is usually made from ceramics, titanium and other metals (3, 7, 45).
Apart from biocompatibility, the major work on axial blood pump development has recently been focused on reducing pump size and improving pump efficiency while reducing hemolysis and thrombosis. A wide variety of tools have been applied to help investigate and optimize the performance of axial blood pumps prior to in-vivo trials. Since the 1990's numerical simulation of fluid dynamics has steadily improved . Computational fluid dynamics (CFD) models can be incorporated with geometric pump models and used to predict pump performance . Some aspects of fluid flow, such as shear stress, are difficult to measure directly but can easily be calculated from CFD models [30,31,32]. The challenge with CFD modeling of blood pumps is that, blood is a multi-component fluid, resulting in very non-ideal behavior. Specifically the viscosity of blood, a parameter directly influencing shear stress and hemolysis, is nonlinear as the various components such as red blood cells are elastic and tend to agglomerate and separate as a function of applied shear. Models of blood behavior have been improving as have models predicting blood trauma. Combining pump CFD models with blood damage models can greatly enhance pump performance and blood damage prediction. This allows extensive optimization of the pump design before physical samples are produced and tested. CFD models are routinely used for optimizing and predicting the effect of various parameters such as the impeller to pump housing radial gap . Larger radial gaps between the impeller and housing result in flow reversal at the impeller blade tips and reduced pumping efficiency. Smaller gaps improve pump performance but result in higher shear rates, and may contribute to hemolysis. Additionally CFD models can be used to identify areas of stagnation of flow inside the pump where thrombosis might occur. Once a pump is produced several techniques can be applied to measure the pumps performance and help correlate the CFD model. Power consumption, flow and pressure can be measured in a fluid test loop . Pump components are often placed in clear housings to allow photographic investigations via various techniques such as particle imaging velocimetry (PIV), particle track velocimetry (PTV), laser Doppler anemometry (LDA) and laser Doppler velocimetry (LDV) [23,31]. LDV and LAD rely of the Doppler shift of laser light reflected from particles in the flow either parallel (LDV) or tangential (LDA) to the direction of the incident laser light. The result is a one component measurement of the velocity of fluid flow at a single point. PIV and PTV can measure the velocity of a flow in two dimensions. Generally a clear test fluid is seeded with small reflective particles. In the PTV technique a single photograph is taken of the target area within the pump. With an appropriate shutter speed the motion of the particles will result in blurred tracks in the resulting image. The length of the track is proportional to the speed of the particle, and the direction can be accessed from the orientation of the track. In PIV the flow is generally illuminated by a sheet of laser light giving a two-dimensional field of illumination. Two images are taken in rapid succession at high shutter speeds. The images are then analyzed to determine the relative displacement of particles from one fame to the next, yielding a twodimensional map of fluid velocity in the plane of illumination. While these optical techniques are extremely valuable they rely on having optical access to the working fluid, necessitating a clear housing or placement of optical access windows in the pump housing. One technique which allows the investigation of flow inside an opaque housing is nuclear magnet resonance imaging (NMR). NMR has the advantage of allowing the use of blood as the working fluid as well. NMR has been applied to measure flows in capillary tubes and will undoubtedly see more application in the investigation of artificial blood pumps in the future [23, 45].
Future Prospects for Axial Flow Blood Pumps
Continuous flow LVADs were initially viewed as a temporary solution until a donor heard could be sourced for transplantation. These LVADs proved to be successful in sustaining the patient in this "bridge to transplant" role, additionally it was noted that the condition of the unloaded heart often improved remarkably. This reverse remodeling has prompted some to subsequently explant the LVAD. While it is clear that the native heart can significantly recover with usage of the LVAD [32, 33, 38], the long-term viability of the patient after device explanation is less certain [8, 32]. With appropriate patient screening and medical treatment it is believed that this role of "bridge to recovery" will increase for axial flow blood pumps in the future [3, 34] As experience with the various LVADs increases and their reliability improves, more people are receiving them for longer and longer durations. Given that the devices can be completely implanted and have a good probability of exceeding 10 year life spans it is highly probably that their role as a destination therapy will expand. It is also very likely that they will be applied to earlier stages of heart failure and to a wider variety of patients including pediatric and emergency cases .
Historically we have had an anthropomorphic bias towards pulsatile artificial hearts. While it is possible that long term non-pulsatile blood flow may have some adverse physiological consequences, this must be balanced against the available options. Axial pumps are simpler, smaller, more easily implantable and more efficient than pulsatile pumps. Given the greater reliability, reduced invasiveness and possibility of infection, and longer life time of axial pumps, it likely that in the long term their benefits greatly outweigh any possible problems associated with their lack of pulse. With the increasing numbers of axial pumps in the field and ever increasing durations of usage they are likely to become more important not just as a bridge to transplant, but as a destination therapy as well. With better understandings of the reverse remodeling, more parents will have access to axial flow blood pumps as a bridge to recovery, allowing eventual device explantation. This is, perhaps, the most exciting possibility as it does not depend on a heart donor, or permanent artificial support. Finally with the small catheterbased axial pumps they are opening a whole new role as a "bridge to decision" temporary support.
[1.] Zareba KM. The Artificial Heart--Past, Present and Future. Medical Science Monitor 2002; 8(3): RA72-77
[2.] Wang W, Zhu DM, Ding WX. Development of Mechanical Circulatory Support Devices in China. Artificial Organs 2009; 33(11):1009-1014
[3.] Sayer G, Naka Y, Jorde UP. Ventricular Assist Device Therapy. Cardiovascular Therapeutics 2009; 27:140-150
[4.] Loisance D. Mechanical Circulatory Support: a Clinical Reality. Asian Cardiovasc Thorac Ann 2008; 16:419-431
[5.] Morris RJ. Total Artificial Heart- Concepts and Clinical Use. Thoracic and Cardiovascular Surgery 2008; 20(3), 247-254
[6.] Boilson BA, Raichlin E, Park SJ, Kushwaha SS. Device Therapy and Cardiac Transplantation for End-Stage Heart Failure. Current Problems in Cardiology 2010; 35(1), 8-64
[7.] Saisho R, Ohshugi T, Watada M, Yong-Jae K, Ohuchi K, et. al. The re-design of Transformer portion in Transcutaneous Energy Transmission System for Left Ventricle Assist Device. World Congress on Medical Physics and Biomedical Engineering, 2006; 1(14), 3193-3196
[8.] Jugdutt BI. Current and Novel Cardiac Support Therapies. Current Heart Failure Reports 2009; 6:19-27
[9.] Chan WK, Yu SCM, Chua LP, Wong YW. Visualization of Relative Flow Patterns in Centrifugal Blood Pumps. Journal of Mechanical Science and Technology, 2001; 15(12), 1869-1875
[10.] Pagani FD, Miller LW. et al. Extended Mechanical Circulatory Support With a Continuous-Flow Left Ventricle Assist Device. Journal of the American College of Cardiology 2009; 54(4):312-321
[11.] Heilmann C. Geisen U, Benk C, Berchtold-Herz M, Trummer G, Schlensak C, Zieger B, Beyersdorf F. Hemolysis in Patients with Ventricular Assist Devices: Major Differences Between Systems. European Journal of Cardio-thoracic Surgery 2009; 36:580-584
[12.] Saito S, Nishinaka T. Chronic nonpulsatile blood flow is compatible with normal endorgan function: implications for LVAD development. J. Atrif. Organs 2005; 8:143-148
[13.] Bourque K, Dague C, Farrar D, Harms K, Tamez D, Cohn W, Tuzun E, Poirier V, Frazier H. In Vivo Assessment of a Rotary Left Ventricular Assist Device-induced Artificial Pulse in the Proximal and Distal Aorta. Artificial Organs 2006; 30(8):638-642
[14.] Seyfarth M,, Sibbing D, Bauer I, Frohlich G, Bott-Flugel L, Byrne R, Dirschinger J, Kastrati A, Schomig A. A Randomized Clinical Trial to Evaluate the Safety and Efficacy of a Percutaneous Left Ventrical Assist Device Versus Intra-Aortic Baloon Pumping for Treatment of Cardiogenic Shock Caused by Myocardial Infarction. Journal of the American College of Cardiology 2008; 52(4):1584-8
[15.] Hussein DH, Gitano-Briggs H, Abdullah MZ. Design Analysis and Performance Prediction of the Cardiac Axial Blood Pump. Research Journal of Biological Sciences 2009; 4(6):637-643
[16.] Lim TM, Zhang D. Development of Lorentz Force-Type Self-bearing Motor for an Alternative Axial Flow Blood Pump Design. Artificial Organs 2006; 30(5):347-353
[17.] Jahanmir S, Hunsberger AZ, Ren Z, Heshmat H, Hshmat C, Tomaszewski MJ, Walton JF. Design of a Small Centrifugal Blood Pump With Magnetic Bearings. Artificial Organs 2009; 33(9):714-726
[18.] Meyer A, Struber M. Chronische Therapie durch links-ventrikulare Unterstutzungssyseme bei terminaler Herzinsuffizienz. Herz 2009; 34:148-153
[19.] Masuzawa T, Shiroh E, Kato T, Okada Y. Magnetically Suspended Centrifugal Blood Pump with an Axially Levitated Motor. Artificial Organs 2003; 27(7), 631-638
[20.] Baldwin TJ, Borovets HS, Duncan BW, Gartner MJ, Jarvik RK, Weiss WJ, Hoke TR. The National Heart, Lung, and Blood Institute Pediatric Circulatory Support Program. Circulation 2006; 113:147-155
[21.] Stone ME. Current Status of Mechanical Circulatory Assistance. Seminars in Cardiothoracic and Vascular Anesthesia 2007; 11(3):185-204
[22.] Morshuis M, El-Banayosy A, Arusoglu L, Koerfer R, Hetzer R, Wieselthaler G, Pavie A, Nojiri C. European experience of Duraheart magnetically levitated centrifugal left ventricular assist system. European Journal of Cardio-thoracic Surgery 2009; 35:1020-1028
[23.] Nojiri C. Implantable Left Ventricular Assist System. Circulation Journal 2009; A:48-54
[24.] Behbahani M, Behr M, Hormes M, Steinseifer U, Arora D, Coronado O, Pasquali M. A Review of Computational Fluid Dynamics Analysis of Blood Pumps. Euro. Jnl. Of Applied Mathematics 2009; 20:363-397
[25.] Bhavsar SS, Kapadia JY, Chopski SG, Throckmorton AL. Intravascular Mechanical Cavopulmonary Assistance for Patients With Failing Fontan Physiology 2009; 33(11)977-987
[26.] Throckmorton AL, Kishore RA. Design of a Protective Cage for an Intravascular Axial Flow Blood Pump to Mechanically Assist the Failing Fontan. Artificial Organs 2009; 33(8):611-621
[27.] Fan HM, Hong FW, Ye L. Design of Implantable Axial-Flow Blood Pump and Numerical Studies on its Performance. Journal of Hydrodynamics 2009, 21(4):445-452
[28.] Wieselthaler GM, Schima H, Zipfer D, Thoma H, Losert U. Fourty years of development, experimental and clinical application of mechanical circulatory support at the Medical University of Vienna. Wien Klinische Wochenschrift 2008; 120:15-20
[29.] Triep M, Brucker C, Schroder, Siess T. Computational Fluid Dynamics ond Digital Particle Image Velocimetry Study of the Flow Through and Optimized Micro-axial Blood Pump. Artificial Organs 2006; 30(5):384-391
[30.] Zhang J, Gellman B, Koert A, Dasse KA, Gilbert RJ, Griffith BP, Wu ZJ. Computational and Experimental Evaluation of the Fluid Dynamics and Hemocompatibility of the CentriMag Blood Pump 2006; 30(3):168-177
[31.] Nugent AH, Bertram CD. Measurement and Analysis of the Flow Field in a Pulsatile Ventricular Assist Device. Journal of Engineering in Medicine 2007; 222:563-571
[32.] Klotz S, Danser AHJ, Burkhoff D. Impact of left ventricular assist device (LVAD) support on the cardiac reverse remodeling process. Progress in Biophysics and Molecular Biology 2008; 97:479-496
[33.] Matsumiya G, Saitoh S, Sakata Y, Sawa Y. Myocardial Recovery by Mechanical Unloading With Left Ventricular Assist System Circulation Journal 2009; (73):1386-1392
[34.] Drakos S, Terrovitis JV, Anastasiou-Nana MI, Nanas JN. Reverse Remodeling during long-term mechanical unloading of the left ventricle. Journal of Molecular and Cellular CARDIOLOGY 2007; 43: 231-242
[35.] Hussein, D., Gitano-Briggs, H. (2008). Outlet Optimization of the Centrifugal Blood Pump. Journal of Engineering and Applied Sciences. 3, 702-707.
[36.] Hehir, D., Niebler, R., Brabant, C., Tweddell, J., Ghanayem, N. (2012). Intensive Care of the Pediatric Ventricular Assist Device Patient. World Journal for Pediatric and Congenital Heart Surgery. 3, 58-66.
[37.] Throckmorton AL, Downs EA, Hazelwood JA, Monroe JO, & Chopski SG. (2012). Twisted cardiovascular cages for intravascular axial flow blood pumps to support the Fontan physiology. The International Journal of Artificial Organs. 35, 369-75.
[38.] Downs EA, Moskowitz WB, Throckmorton AL. (2012). Steady flow analysis of mechanical cavopulmonary assistance in MRI-derived patient-specific fontan configurations. Artificial Organs. 36, 972-80.
[39.] Throckmorton AL, Kapadia J, Madduri D. (2008). Mechanical axial flow blood pump to support cavopulmonary circulation. The International Journal of Artificial Organs. 31, 970-82.
[40.] Hussein, D.H., Gitano-Briggs, H., Addullah, M.Z. (2009). Design Analysis and Performance Prediction of the Cardiac Axial Blood Pump. ASME Journal of Medical Devices. 4/6, 637-643.
[41.] Sciolino, M.G. (2012). Development, optimization, and twisted adjustment of an axial flow blood pump for Fontan patients. Richmond, Va, Virginia Commonwealth University. http://hdl.handle.net/10156/4120.
[42.] Chopski, S.G. (2010). Particle image velocimetry measurements of the total cavopulmonary connection with circulatory flow augmentation. Richmond, Va, Virginia Commonwealth University. http://hdl.handle.net/10156/2863.
[43.] Kapadia, J.Y. (2009). Development of a mechanical cavopulmonary assist device for the failing fontan patients. Richmond, Va, Virginia Commonwealth University. http://hdl.handle.net/10156/2667.
[44.] Sciolino, M.G. (2012). Development, optimization, and twisted adjustment of an axial flow blood pump for Fontan patients. Richmond, VA, Virginia Commonwealth University. http://hdl.handle.net/10156/4120.
[45.] Swalen, Marcel Johannes Petrus. (2012). Study of a bi-directional axial flow blood pump. Brunel University School of Engineering and Design PhD Theses. http://bura.brunel.ac.uk/handle/2438/6343.
[46.] Hussein, D.H., Gitano-Briggs, H., Addullah, M.Z. (2009). Cardiac Axial Blood PUmp Analysis and Performance Prediction. ASME Journal of Medical Devices. 3, 027549.
Dhyaa Kafagy (1), Horizon Gitano-Briggs (2),
(1) Virginia Commonwealth University, Richmond, Virginia, USA.
(2) University of Kuala Lumpur, Malaysian Spanish Institute, Malaysia.
Corresponding author: Prof. Dhyaa Kafagy, Dhyaa.Kafagy@gmail.com
Received 1 July 2012; Accepted 7 July 2013; Available online 21 July 2013
Table 1: Relative comparison of various parameters from three types of blood pumps (2,11,25,26,27) Pulsatile Pump Centrifugal Axial Flow Speed ~ 60 bps 2000-4000 RPM 6000-12000 RPM Size Large Medium Small Hemolysis (NIH) High 0.0007 - 0.009 0.0028 - 0.03 Thrombosis Low Medium High
|Printer friendly Cite/link Email Feedback|
|Author:||Kafagy, Dhyaa; Gitano-Briggs, Horizon|
|Publication:||Trends in Biomaterials and Artificial Organs|
|Date:||Jul 1, 2013|
|Previous Article:||Polyvinyl siloxanes in dentistry: an overview.|
|Next Article:||Role of magnets in orthodontics and dentofacial orthopedics: a comprehensive review.|