CA2206644A1 - Ventricular assist device comprising enclosed-impeller axial flow blood pump - Google Patents

Ventricular assist device comprising enclosed-impeller axial flow blood pump

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Publication number
CA2206644A1
CA2206644A1 CA002206644A CA2206644A CA2206644A1 CA 2206644 A1 CA2206644 A1 CA 2206644A1 CA 002206644 A CA002206644 A CA 002206644A CA 2206644 A CA2206644 A CA 2206644A CA 2206644 A1 CA2206644 A1 CA 2206644A1
Authority
CA
Canada
Prior art keywords
pump
blood
vad
axial flow
rotor
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Abandoned
Application number
CA002206644A
Other languages
French (fr)
Inventor
L. Conrad Pelletier
Alain Girard
Andre Garon
Rosaire Mongrain
Michel Carrier
Stephane Trudelle
Ricardo Camarero
Original Assignee
L. Conrad Pelletier
Alain Girard
Andre Garon
Rosaire Mongrain
Michel Carrier
Stephane Trudelle
Ricardo Camarero
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by L. Conrad Pelletier, Alain Girard, Andre Garon, Rosaire Mongrain, Michel Carrier, Stephane Trudelle, Ricardo Camarero filed Critical L. Conrad Pelletier
Priority to CA002206644A priority Critical patent/CA2206644A1/en
Priority claimed from CA002292432A external-priority patent/CA2292432A1/en
Publication of CA2206644A1 publication Critical patent/CA2206644A1/en
Application status is Abandoned legal-status Critical

Links

Classifications

    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/101Non-positive displacement pumps, e.g. impeller, centrifugal, vane pumps
    • A61M1/1029Drive systems therefor
    • A61M1/1031Drive systems therefor using a motor with canned rotor, i.e. a motor enclosed within a casing along with the rotor so that the motor bearings are lubricated by the blood that is being pumped
    • A61M1/1036Drive systems therefor using a motor with canned rotor, i.e. a motor enclosed within a casing along with the rotor so that the motor bearings are lubricated by the blood that is being pumped using rotating magnets for driving
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/1008Tubes; Connections therefor
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/101Non-positive displacement pumps, e.g. impeller, centrifugal, vane pumps
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/101Non-positive displacement pumps, e.g. impeller, centrifugal, vane pumps
    • A61M1/1012Constructional features thereof
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/101Non-positive displacement pumps, e.g. impeller, centrifugal, vane pumps
    • A61M1/1029Drive systems therefor
    • A61M1/1031Drive systems therefor using a motor with canned rotor, i.e. a motor enclosed within a casing along with the rotor so that the motor bearings are lubricated by the blood that is being pumped
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/12Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps implantable into the body
    • A61M1/122Heart assist devices, i.e. for assisting an ailing heart, using additional pumping means in the blood circuit
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/12Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps implantable into the body
    • A61M1/125Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps implantable into the body intravascular, i.e. introduced or implanted in an existing blood vessel
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M1/00Suction or pumping devices for medical purposes; Devices for carrying-off, for treatment of, or for carrying-over, body-liquids; Drainage systems
    • A61M1/10Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps
    • A61M1/12Blood pumps; Artificial hearts; Devices for mechanical circulatory assistance, e.g. intra-aortic balloon pumps implantable into the body
    • A61M1/127Energy supply devices, converters therefor
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61MDEVICES FOR INTRODUCING MEDIA INTO, OR ONTO, THE BODY; DEVICES FOR TRANSDUCING BODY MEDIA OR FOR TAKING MEDIA FROM THE BODY; DEVICES FOR PRODUCING OR ENDING SLEEP OR STUPOR
    • A61M2205/00General characteristics of the apparatus
    • A61M2205/35Communication
    • A61M2205/3507Communication with implanted devices, e.g. external control
    • A61M2205/3523Communication with implanted devices, e.g. external control using telemetric means

Abstract

The present invention relates to a Ventricular Assist Device (VAD) comprising an axial flow blood pump which can be used as bridge to heart transplantation and as a permanent implant. The pump of the present invention is of an enclosed-impeller axial flow design, in which the rotor comprises impeller blades that intersect at an axis of rotation to form a hub of a relatively small size. The minimization of the hub volume allows a larger internal volume of the rotor to be occupied by the fluid. The pump can also be viewed as a double casing axial flow blood pump in which the permanent magnets required for the DC motor can be embedded in the enclosing tube (rotor) and the motor windings can be enclosed in the pump casing (stator). The configuration thereby optimizes the electrical coupling and this minimizes power requirement.
In a preferred embodiment, the axial flow blood pump of the present invention is implantable in the ventricle of a patient. In such an intra-ventricular VAD, although the pump has a continuous speed, the pump output is pulsatile when inserted in the native heart.

Description

CA 02206644 1997-0~-30 TITLE OF THE INVENTION
VENTRICULAR ASSIST DEVICE COMPRISING AN
ENCLOSED-IMPELLER AXIAL FLOW BLOOD PUMP

This invention relates to ventricular assist devices and more particularly to enclosed-impeller axial flow blood pumps.

BACKGROUND OF THE INVENTION
In North America, heart related diseases are still the leading causes of death. Among the causes of heart mortality are congestive heart failure, cardiomyopathy and cardiogenic shock. The incidence of congestive heart failure increases dramatically for people over 45 years. In addition, a large part of the population in North America is now entering this group of age. Thus, the number of people who will need treatment for these types of diseases will reach a larger proportion of the pop~ ion. Many complications related to congestive heart failure including death could be avoided and many years of life could be saved if proper treatments were available.
The type of treatments for heart failure depends on the extent and severity of the illness. Many patients can be cured with rest and drug therapy but there is still severe cases that require heart surgery including heart transplantation. Actually, the mortality rate for patients with cardiomyopathy who received drug therapy is about 25 % within 2 years and there still is some form of theses diseases that cannot be treated medically. One of the last options that remain for these patients CA 02206644 1997-0~-30 is heart transplantation. Unfortunately, according to the procurement agency UNOS (United Network for Organ Sharing in United States) the waiting list for heart transplantation grows more than 2 times faster than the number of heart donors.
Considering these arguments, it appears imperative to offer alle" ~dti~e l~ eatmenls to heart l~ ansplantation. The treatment should not only provide life extent but also improve quality of life. In this context, mechanical circulatory support by means of Ventricular Assist Devices (VAD) is a worthwhile possibility in front of a large deficiency in the number of available organ donors. In the eighties, successful experiments with mechanical hearts and VAD serving as bridge to transplantation increased significantly. The accumulated knowledge in all aspects of patients care, device designs and related problems led to the use of VAD as permanent implants. Now, it appears appropriate to address the problem of end stage heart failure with permanent mechanical heart i",,~la"~s. Among the various mechanical supports, axial flow VAD which aims at durability of 5 to 10 years is a very interesting approa~ ,. It is estimated that 2 000 patients per year in Canada and 30 000 patients per year in the United States could benefit from VAD.
In 1980, the National Heart, Lung and Blood Institute (NHLBI) in United-States defined the characteristics for an implantable VAD (Altieri and Watson, 1987). These characteristics include medical requirements that are to restore hemodynamic function (pressure and cardiac index) avoid hemolysis, prevent clot formation infection and bleeding, and minimize anti-coagulation requirement. More technical characteristics include: small size, control mode, long durability ( > 2 years), low heating, noise and vibration.

CA 02206644 1997-0~-30 Ventricular Assist Devices (VAD) can be used in several circumstances where a patient has poor hemodynamic functions (low cardiac output, low ejection fraction, low systolic pressure). Whatever the origin of the cardiac failure, the goal of the VAD is to help the heart in his 5 pumping action. The VAD unload the heart by producing an enhanced circulation and thus restoring the hemodynamic functions which will provide good end organ perfusion. Many devices can achieve these goals, however they are not optimal, hemolysis and thrombus formation are still important problems to investigated.
In the 70s, the first approach to solve the problem of mechanical support was to imitate as much as possible the heart physiology. This resulted in the development of several pulsatile devices, some of these initial designs are still used. The first developments were pneumatically driven devices while a second generation of pumps was 15 electrically ~tu~ted. In the 90s, a new generation of pumps has emerged which ad-JIesses certain problems ~ssoci~ted with previous devices (size and power consumption). They are non-pulsatile devices mainly divided into two groups: centrifugal and axial flow blood pumps.
What follows is a review of models that are still in use, 20 either for clinical or research purposes.

I. PULSATILE DEVICES
Until now, pulsatile VAD have been the only devices used in the clinical stage for humans although several other designs are 25 being used at the research stage on animals. The four more popular devices are from: Novacor Inc., Thermo Cardiosystem Inc. (TCI), Thoratec Inc. and CardioWest Inc. All these devices have been widely CA 02206644 1997-0~-30 used in human and have been approved by the FDA (Food and Drug Administration).

1. Pneumatic 5 1.1 Tl-~ratec Inc.
The first pneumatical device which produced good clinical outcome was the Pierce-Donachy VAD. The first trial in human began in 1976. The device now owned by Thoratec Laboratory, Berkeley, CA is simple, versatile and allows to support left or left and right 1 0 ventricles.
The Tl ,oralec VAD is a sac type prosthetic ventricle. The polyurethane elastomer sac is contained in a machined polycarbonate housing. Two mechanical Bjork-Shiley valves are used at the inlet and outlet to allow unidirectional blood flow. A port is connected to the driving 15 console which provides pressurized air contained in two bottles of gas.
As air enters the VAD, the blood sac is co,l,pressed and blood is ejected via the outlet. Then a negative pressure is created which facilitates the VAD filling. Figure 1 shows a sketch of the Pierce-Donachy VAD.
The console allows three operating modes. The first is 20 asynchronous with a fixed rate set at the console. The two other modes are variable. In the second mode, a Hall-effect sensor signals to the console when the blood sac is full which in turns initiates the emptying process. In this mode, the sac fills faster or slower according to the changes in preload, changing the rate in consequence. Finally, in the 25 third mode, the emptying pr~cess is synchronized with the R-wave of the ECG. For this device, the maximum volume stroke is 65 mL and allows a flow of 5.5 Umin.

CA 02206644 1997-0~-30 ;

This VAD is surgically placed in a para-corporal setting;
the two blood ports exit via the abdomen and the VAD is laying over it.
The inlet is usually connected to the left atria and the outlet is connect to the ascending aorta. The two flow cannula are tunnelled under the skin 5 from the heart to the exit site. The first trials were done on patients that could not be weaned from cardiopulmonary bypass after heart surgery (Copeland et al. 1985).

1.2 CardioWest Inc.
10In the same period, several groups were working on the development of similar devices, including the group of Symbion lead by Dr. Robert Jarvik. The principle of their device is almost identical to the one of Pierce-Donachy except that it is a Total Artificial Heart (TAH).
The first patient to receive successfully the device as a 15permanent implant was Barney Clark in 1982, who survived 112 days.
The Jarvik-7 is now sold by CardioWest, Tucson, AZ and is used only as a bridge to transplantation.
This TAH consists of two ventricles with two Medtronic-Hall valves on each of them. The semi rigid housing is made of 20 polyurethane polymer and contains an inflatable diaphragm. This diaphragm is composed of four layers of Biomer with graphite powder between them for rupture prevention. Two air drive lines, one on each ventricle, produce diaphragm displacement and heart emptying. These lines exit above the umbilicus and are connected to a console. The 25 console allows several setlings: heart rate, systolic pressure, % of systole and finally negative pressure to facilitate filling of the device. Since the heart has been excised, no gating is needed. When the patient is doing CA 02206644 1997-0~-30 exercise, the heart rate and the other controls always need to be manually adjusted. The device has a maximum filling volume of 70 mL.
When adding to this the supplementary volume needed for the housing and the graft for the anastomoses the assembly results in a relatively 5 large device. The resulting volume limits its implantation to patients weighing at least 79 kg and measuring not less than 1.75 m. These size requirements exclude children, small men and most women.

1.3 Thermo Cardio Inc.
Another pulsatile device widely used is the HeartMate Implantable Pneumatic LVAD (IP-LVAD) by TCI (Thermo Cardiosystem Inc., Woburn, MA). Clinical trials began in 1986 with a device designed for long term use (longer than two years). The blood pump has a titanium alloy housing with two chambers (air and blood) separated by a flexible 15 polyurethane diaphragm (Biomer) Iying on a pusher-plate (Figure 2). The inlet and outlet of the pump are equipped with porcine valves. The pump housing measures 11.2 cm in diameter and 4.0 cm in thickness.
The surfaces in contact with the blood are specially designed to facilitate cell deposition resulting in a biological coating for 20 thrombogenesis minimization. (Timothy et al., 1995).
As in other pneumatic devices, blood ejection is obtained by pressurized air that moves the pusher-plate and the diaphragm. For the filling, the pressure is brought back to a zero level then the VAD is passively filled. The maximum output stroke is about 83 25 mL. The control console of the IP-LVAD provides the same three control modes as with the Thoratec VAD.

CA 02206644 1997-0~-30 For the implantation, the pump is positioned intra-abdominally or preperiloneally just below the diaphragm. Pump inflow p~-sses through the diaphragm and is connected to the apex. The pump outflow graft is anastomosed to the ascending aorta. The percutaneous 5 air drive line is tunnelled and exits by the left lateral side above the iliac crest.
The main disadvantage of these pneumatic devices is that the patient is always tethered by a 10 foot drive line to a console generally large and heavy (about 100 kg and the size of an household 10 stove, a smaller portable console is under evaluation). This greatly reduces the mobility of the patients. To overcome this problem, more recent LVAD are electrically powered. These types of devices are briefly described in the following section.

15 2. Electromechanical The main advantage of electrical powered VAD is the increase in mobility achieved by a smaller controller made portable and a set of rechargeable batteries which untethered the patient and enable him to walk freely. The counterpart is that more mechanical components 20 are required in the device which makes its development more complex.
Pulsatile VAD always contain a blood and an air cha"lber. In elect,ical pulsatile devices, the air chamber usually contains the motor or the ~ctl~tor that pushes against the diaphragm. When the diaphragm is moved to eject blood from the blood chamber, an equal 25 volume of air has to enter the air chamber. This air displacement can be achieved in two ways: with a percutaneous vent connected to the air chamber and with a compliance chamber. The latter is a small bag CA 02206644 1997-0~-30 connected to the air chamber of the VAD so that air moves from the bag (deflation) to the pump in the ejection phase and back to the bag while the blood is filling (bag inflation). This solution is not favored because it has shown several practical complication ex. the bag often needs to be 5 refilled due to air diffusion through the bag membrane.

2.1 Thermo Cardio Inc.
The TCI company has also developed a VE-LVAD
(Vented Electric LVAD). This pump is almost identical to the IP-LVAD
10 except for its actuation process. It has the same housing, valve, volume stroke and surgical implantation. For their electric LVAD, the non blood charnber contains a small motor. As the motor turns on, it pushes two ball bearings on helical cams which move the pusher-plate and pressurize the blood chamber thus ejecting the blood from the pump.
Only two control modes are available for the VE-LVAD
compared to three with the IP-LVAD (no ECG gating is available).
Concerning the surgical i,nplantalion, the only difference is that there are two exit sites, one for the electrical cable and one for the pump vent.

20 2.2 Novacor Inc.
The VAD of Novacor Inc. (Oakland, CA) a division of Baxter Healthcare col ~oralion contains a polyurethane sac that contains the blood. The sac is bounded by two pusher-plates. The actuation is made by a solenoid and a spring which bring closer and farther apart the 25 plates and consequently empty and fill the LVAD. Two bovine valves are also included to provide unidirectional blood flow.

CA 02206644 1997-0~-30 The large device weighting about 1 Kg is placed in the abdomen in the same way as the HeartMate LVAD. However, there is only one exit site, the electrical cable exits by the percutaneous vent. The VAD is connected to a compact controller worn on a belt. Two 5 rechargeable batteries are also provided, a primary one and a second one used as backup. The first clinical trial with this LVAD has begun in 1 984.

3. Conclusion Since 1984, 1286 cases of peoples receiving one of the devices described above were registered at the American Society for Artificial Internal Organs. Table 1 summarizes some of their characteristics (Arabia et al., 1996). The HeartMate VAD is the most used with 39 % of market with a large majority of pneumatic devices. The best 15 sl ~ccess rate is observed with the Novacor system although most systems shows comparable results. The success rate is defined as the percenlage of patients who sllccessfully underwent a heart transplant among the people who sllccessfully received the mechanical support.

CA 02206644 1997-0~-30 Table 1: Technology, sl~ccess rate and market share for currently widespread devices.
Market Success Company Support Cor", ~ller Actuator Share Rate (%) (%) Novacor LVAD Console Electro-mechanic 14 90 LVAD Portable Electro-mechanic 14 92 TCI LVAD Console Pneumatic 35 89 HeartMate LVAD Portable Electric 4 89 Thoratec LVAD Console Pneumatic 10 93 BiVAD Console Pneumatic 19 81 CardioWest TAH Portable Pneumatic 4 92 As reported, although pulsatile devices have been mostly used, many illlpGilalll problems are still ~ssoci~ted therewith.
Their large sizes limit their uses in corpulent men and their mechanical 15 complexities make their uses hazardous for periods longer than two years. In addition, the hemolysis rate remain high with these pumps.
Moreover, the presence of the transcutaneous vent is an important risk of infection.
Consequently, a new generation of blood pump based 20 on non-pulsatile ,> principles have been designed to minimize problems associated with pulsatile ventricular assist devices. These pumps are described in the next sections.

Il. NON-PULSATILE DEVICES
Non-pulsatile devices are divided into two main categories, centrifugal pumps and axial flow blood pumps, both of which are described below.

CA 02206644 1997-0~-30 In non-pulsatile VAD, an impeller is enclosed in a housing and the continuous rotation of the impeller produces the pumping action. The faster the rotation, the higher the blood flow. These devices are called non-pulsatile or continuous becAuse they operate at constant speed. Most axial flow blood pumps operate around 10 000 rpm (Rolalions Per Minute). However, in in-vivo conditions, there is a dynamic range (about 1000 rpm around the operating point) over which the output flow is pulsatile. Since the native heart is still contracting, a pressure difference between the ventricle (inlet) and the aorta (outlet) is created.
This pressure variation will produce a variation in the pump flow. The range of rpm over which pulsatile flow occurs is small; at lower speed back flow is observed (in diastole) and at higher speed the heart is completely unloaded. In the latter case, no pressure variation occurs resulting in non-pulsatile flow.
Many advantages are associated with the use of non-pulsatile devices and they all have a strong impact on the physiology as well as on the clinical management. These are:

Size - Non-pulsatile devices are much smaller in volume than pulsatile ones, around 25 cc for axial pump, and 100 cc for centrifugal pumps compared to 150 cc and more for pulsatile devices. For comparison a complete axial-flow VAD is usually smaller than the graft used for pulsatile pump. The clinical impact is the possibility to use this type of VAD in small adults as well as in children. Also the small dimensions enable to place the pump in a more orthotopic position; that is, in the thorax near the heart instead of the upper abdomen. This eliminates the use of long cannula passing through the diaphragm, thus reducing the CA 02206644 1997-0~-30 risk of rejection. FU, ll ,er",ore, for axial flow VAD, the shape and size can be made to place the VAD in an intra-ventricular position.

Power - The electrical power required to drive a continuous device is less 5 than for pulsatile devices. Indeed, when a motor is turned on, it is in transient state until it takes its normal speed, afterwards it is in steady state. The transient state produces a peak in power demand and the steady state requires much less power. Since pulsatile devices constantly work in transient state, their power requirements are much more 10 important than for continuous devices. In fact, the continuous VAD
requires about 30 to 40 % less power than pulsatile VAD. This means that smaller batteries can be used or longer autonomy can be achieved with the same bdtleries. Another significant consequence is that there is less heat production which could be potentially damaging for the 15 surrounding tissues.

Simplicity - Non-pulsatile VAD are mechanically simpler than pulsatile VAD in that they do not require any valve, diaphragm, blood sac, vent or compliance chamber. Continuous VAD is made of a simple motor to 20 which is coupled an impeller contained in an housing. One important advantage of a simple mechanical design is that it enables better durability. Durability of 5 to 10 years (Nosé, 1995 ed., ~larvik, 1995) could be achieved with continuous VAD compared with two years for pulsatile VAD (Pierce, 1996). In principle, this would allow not only to use it as a 25 bridge to transplantation but also as long term mechanical support.
Among the advantages of a simpler design is a lower production cost.

CA 02206644 1997-0~-30 This implies that a larger diffusion could be achieved hence a greater number of patients could benefit from the technology.

He".oly~is - I len,olysis or tearing of the red blood cells can be estimated 5 with a parameter called the Normalized Index of Hemolysis (NIH). The NIH of most pulsatile devices is around 0.04 (Nosé, 1995 ed.) whereas for continuous devices the NIH drops to a range between 0.002 to 0.004, that is about 10 to 20 times smaller. As a consequence, these devices are less traumatic for blood. Pulsatile VAD has a higher NIH because of 10 the high dP/dt (pressure variation) required to produce pulsatile flow;
also, the presence of two valves in the case of mechanical valves may also co,)l, ibute to increase the hemolysis.

Infection - Interestingly, probability of infection with continuous devices 15 is red~ ~ce-i Since the transcutaneous vent of pulsatile devices is an open door for opportunist infections daily cleaning is required. Moreover, the pos~ ility to place the VAD in the thorax rather than in the abdomen also contributes to reduce the risk of infection.

20 Patient considerations - Non-pulsatile devices require less maintenance allowing the patient a greater autonomy. Also, as it is now known, most patients with a VAD are discharged from the hospital and returned to a normal life after about a month. Actually, because of the vent in pulsatile devices, patients cannot take a bath or swim since water could enter the 25 motorcompa,l",ent. Continuous VAD are less restrictive and allow more activities for the patient.

CA 02206644 1997-0~-30 1. C~. Itl iruqal Pumps Centrifugal blood pumps were first used in cardio-pulmonary bypass for heart surgery. Based on results obtained with the Bio-Medicus pump (Medtronic Bio-Medicus Inc., Eden Prarie, MN), several groups decided to develop much smaller centrifugal pumps so that they could be totally implantable. In this section, we will not discuss specific characteristics of the design but rather provide a general discussion of the concepts.
Figure 3 illustrates the head of a centrifugal pump. For this type of pump, the rotation of the impeller produces a centrifugal force that drags blood from the inlet port on top to the outlet port at the bottom side. To produce the rotation of the impeller, the head is coupled to a motor. The coupling is made either magnetically by means of permanent magnet located under the impeller and on the rotor or by a shaft located between the impeller and the rotor. Magnetically coupled devices generally show better functionality because no seal is required between the motor and the impeller shaft.
O,ulil~ dlion of this type of pump is concerned with the position and orientation of the inlet port, the clearance between the impeller and the housing and the shape of the blade, all of these design characteristics having a direct impact on the hemolysis.
Among the groups working on this type of devices, there are the Baylors College of Medicine with their C-gyro pump (sponsored by the gove"""enl of Japan), the Vienna University, since the beginning of the 80's (Schima et al., 1994), and a few other groups mainly in Japan.

CA 02206644 1997-0~-30 These devices involved in their basic design the well known technologies of the Bio-Medicus pump recognized for its low hemolysis rate. But to be implantable, the centrifugal pump needed to be much smaller which resulted in an increased of the centrifugal forces.
5 Also to keep an appro,criate blood flow, the rotational speed of these pumps had to be increased or blades added as illustl;3lecl in Figure 3 (the BioMedicus pump is bl~d~less). All of these modifications will col,l,ibute to increase the hemolysis and as a consequence, an hemolysis index similar to the one observed with the Bio-Medicus pump can hardly be 10 achieved. Another problem with centrifugal pumps is that although they are much smaller than pulsatile pumps, they are still too large to be totally implanted in a human thorax and thus eliminating an intra-ventricular implantation.
It is also known that shear stress, the main factor 15 responsible for hemolysis production, is related to friction forces and the exposure time to these forces (Pinotti and Rosa, 1995). In the case of centrifugal pumps, it is difficult to evaluate the duration of blood cells exposure to these forces. As we can see on Figure 4, when new blood enters the pump, a fraction is rapidly ejected but an other one is 20 recira~ ed and will be ejected later. In consequence, blood cells may be exposed for long periods of time which in principle would increase the hemolysis associated with this design.
To overcome the problems with centrifugal pumps, a new type of pumps are now under development. These are axial flow 25 ventricular assist devices. These pumps can decrease the hemolysis rate by having a shorter exposure time to friction forces and by reducing the CA 02206644 1997-0~-30 i"tensity of these forces. Another interesting advantage is that axial flow blood pumps are generally much smaller than centrifugal pumps.

2. Axial flow blood pumps The first commercially available axial flow blood pump was the Hemopump (Medtronic Inc. Minneapolis, MN) used as short term circulatory support. Based on the good results obtained with this device, several groups have initiated the dcvelop,nent of totally implantable axial flow VAD for long term uses.
Four axial flow ventricular assist devices are presently under development. 1 -DeBakey/NASA by the Baylors College of Medicine in conjunction with the NASA Johnson Space Center, 2 - the Jarvik 2000, by Jarvik Research Inc. (New York, NY) and Transicoil Inc.
(Valley Forge, PA), 3 - AxiPump, developed by Nimbus Inc. (Rancho Cordova, CA) in collaboration with the Schools of Medicine and Engineering from the University of Pittsburgh, 4 - The Heart Institute of Japan, Waseda University and Sun Medical Research. All of these groups have already started in-vivo experiments on animals although they still perform in-vitro trials.
Figure 5 gives a schematic of an axial flow blood pump.
The pump consists of a casing which encloses the pump elements. The rotor, located in the center produces the pumping action. The rotor contains permanent magnets; motor windings are located outside the casing. On each side of the rotor, there are two stators. The inflow stator converts laminar flow into a rotational flow similarly the oufflow stator converts the rotational flow into a laminar flow.

CA 02206644 1997-0~-30 An important issue about these types of pumps is the bea, ing system that supports the impeller. The impeller must rotate easily with minimal wear and resistance. Three technologies are available (Jarvik, 1995) to support the impeller: blood-immersed bearing (Jarvik 5 2000 and DeBakey/NASA), sealed bearing (Japan Heart Institute and Nimbus) and magnetically suspended rotor. Actually, the latter is being stu~ ed by a few groups but the technology is still at the research stage.
Concerning the blood-immersed bearing, there is a tendency for ll " ombus formation around the bearing localiol1. To minimize this adverse 10 effect, high speed flow has to be provided in order to produce a good bearing washout. For the case of sealed bearing, there is usually U,ro",bus fo""dlion at the interface of seal and contacting surface. Also, a purge port is required, through which a sterile solution is injected at continuous pressure and flow to ensure that blood elements do not enter 15 the sealed area.

2.1 Jarvik 2000 The Jarvik 2000 is one of the simplest axial flow VAD;
it is illustrated in Figure 6 (on this drawing, blood flows from right to left).20 The pump contains two stators, one at the inflow and one at the oufflow.
They have two functions: they support the bearing shaft around which the impeller will rotate (middle part) and they modify the blood path. The inflow stator initiates the rotation of the flow so that the blade tip of the rotor does not create too much shear stress on blood cell. The oufflow 25 stator is a flow straightener so that blood enters the blood stream with an axial profile. Permanent magnets are enclosed in the center of the impeller and two motor windings are located in the casing on each side CA 02206644 1997-0~-30 of the rotor. This configuration consists of a DC brushless motor this is a simple and durable motor that minimizes the number of mechanical parts. The power cable is connected directly to the DC brushless motor cor,l,oller to cha,)ge motor speed. The device is made of a titanium alloy and measures 5 cm long by 1.8 cm in diameter the displacement volume is 12 mL (technical specifications are summarized in Table 2).
To implant the device the chest is opened by mean of a left thoracotomy and no cardiopulmonary bypass is used. The pump oufflow is anastomosed to the aorta with a Dacron graft. Then a ventriculotomy is made to insert the pump into the ventricle through a sewing ring attached to the apex.
Experiments with this device were conducted on 7 sheep (Kaplon et al. 1996). Animals survived from 3 to 123 days most of the e~ri,nents ended be~use of broken power cable. A rise from 10 to 17 mg/dL of plas",a free hemoglobin was observed but was statistically insiy"iricant. At autopsy no animals showed end-organ dysfunction and no emboli were found. Ongoing experiments with this device appears promising.

2.2 DeBakeY/NASA
The DeBakey/NASA ventricular assist device is very similar to the Jarvik-2000 design. The principle is the same it is based on a DC brushless motor with blood-immersed bearings an impeller in the middle and two fixed parts on each side. Figure 7 illustrates the pump assembly. It is also taught in US Patent 5 527 159.
Blood enters by the left side and passes through a flow straightener that prevents its prerotation then it reaches the CA 02206644 1997-0~-30 induce,/i",peller; the inducer is used to initiate rotation of blood before it reaches the impeller. It is the latter that produces the effective pumping action. Finally a flow diffuser converts the tangential flow into an axial flow. The inducer contains three blades and the i,npeller holds six blades.
5 The three inducer blades are continuous with three of the impeller blades. Each blade of the impeller contains eight cylindrical permanent magnets. Finally a winding is placed outside the pump to complete the motor assembly. The pump is 7.5 cm long and the flow cylinder has a 1.3 cm diameter the larger diameter at the motor winding is 2.7 cm. The 10 casing is made of titanium and the central part is made of polycarbonate.
Polycarbonate will eventually be changed to achieve better biocompatibility. The rotor is held by two bearings the front one being a ~i~conia ball in a sapphire cup and the other a zirconia shaft in a sapphire olive-endstone. See Table 2 for technical specifications.
A few short term animal experiments on calves were conducted using this pump. The pump was in a paracorporal position for experi,nenlal setup convenience. The inflow was inserted in the ventricle through the apex and oufflow was anastomosed to the thoracic aorta.
Although plasma free hemoglobin was in a quite acceptable range 20 thrombus formation occurred at the flow straightener and at the bearing locations (Mizuguchi et al 1995; Kawahito et al. 1996). However the pump has showed good ~p~h ' ty to maintain satisfactory hemodynamic functions for a period of nine days.

25 2.3 Japan Heart Institute The axial flow blood pump developed by this group is quite dirraren~ from the two previous ones ( see Figure 8). It is a sealed-CA 02206644 1997-0~-30 bearing type pump whose motor is separated from the impeller. This is also called a dry motor configuration compared to the Jarvik 2000 and DeBakey/NASA pump which are called wet motor configuration. In the figure, the motor is on the left most side. The impeller is coupled to the motor by a driving shaft. Blood enters the pump by four holes at its basis, p~.sses through the impeller and the flow straightener then exits the pump by the tapered end. The pump is designed so that there is a gradual increase of pressure along the axis.
The surgical placement of the pump is quite different from the other VAD. This pump is the easiest one to install; once the chest is opened only 5 minutes are required to install the pump. An incisio" is made at the apex and the pump is inserted in the left ventricle until the basis of the motor is in contact with the apex. The motor stays outside the heart and is sutured to the heart. The outlet port is passed through the aortic valve and since blood ejects directly in the aorta, no oufflow cannula is required.
The main problem with this VAD is the presence of the seal that requires continuous infusion of sterile solution in the sealed area to prevent blood from entering; 576 mL of solution is required for 24 h of operation. This large volume implies to use an external refillable bag that contains the liquid and a transcutaneous cannula connected to the pump purge port.

2.4 AxiPump This pump is similar to the DeBakey/NASA and the Jarvik 2000 pumps. It is placed beside the heart and connected between the apex and the aorta. It is also a sealed-bearing type pump and so CA 02206644 1997-0~-30 needs a purge system. This system requires a second pump that pushes a 15 mUday of sterile solution in the sealed area. A pump without purge system is now under development. Three animals have been supported for 1 month with the AxiPump (Konishi, 1 996b).

3. Conclusion Table 2 presents characteristics of the four axial flow blood pumps described above. The lower the NIH, the lower the hemolysis. For co",paris~n purpose, the traditional roller-pump has a NIH
of about 0.06 g/100L (Damm et al., 1993), a pulsatile VAD around 0.04 g/100L (Nosé, 1995a) and the HemoPump is between 0.035 to 0.082 g/1 OOL (Mi~ug~ ~chi et al., 1 994b).

Table 2: Technical characteristics of axial flow VAD.
neR~k~y / Japan Heart Jarvik 2000 AxiPump NASA Institute Size (cm) 5.0 X 1.8 7.5 X (1.3 - 2.7) 10 X 1.4 1.4 diameter t~isplacement volume 12 mL 15 mL 62 mL
Weight 45 9 53 9 170 9 205 9 Speed (rpm), 19000 10800 9000 11200 5 L ~ 100 mmHg Power 10W 11 W 8W 6-10W
NIH (g/100L) 0.0029 0.005 0.0045 PFHb (mgldL) 17 17 10 15 All the axial flow ventricular assist devices described above are very attractive but they all have limitations: For the Nimbus and the Japan Heart Institute pumps, the sealed bearing approach requires CA 02206644 1997-0~-30 a purge system and a transcutaneous fluid delivery system. Also, there is ll ,ro,nbus formation at the seal locations. For the Jarvik 2000 and the DeBakey/NASA pumps, there is poor bearing washout which results in thrombus formation at the bearing locations.
Other axial flow blood pumps have also been proposed.
Isaacson in US patent 5,205,721 teaches an axial flow blood pump having an hydrodynamically suspended rotor centrally positioned with respect to the stator. Hydrodynamic bearings are created by two spaces in which blood must flow to create the hydrodynamic support which in tums is associated with shearing forces on the blood. Is~cson teaches three types of impeller blades which significantly add to the volumetric contribution of the hub.
ISAACSOn et al. in US patent 5,211,546 also teaches an axial flow blood pump which is similar to that taught in US 5,205,721.
Indeed, the rotor is suspended radially by hydrodynamic bearings. In certain embodiments, a radially centered thrust bearing element is loc~led within the outlet directed surface of the pump stator inlet section and acts to stabilize the ,c,laliGn of the suspended rotor. Once again, the volumetric contribution of the hub and vanes is significant and the blood flow in such a design is expected to provide significant shearing of the blood.
US patent 5,290,227 of Pasque on the other hand teaches an axial flow blood pump having a rotor assembly described generally as a hollowed-out cylinder provided with rotor vanes which extend to the interior surface of the hollowed cylindrical rotor towards the center axis of the rotor. Fortunately, such a design generates two pumping zones inside the pump. One of these zones is an outer annulus CA 02206644 1997-0~-30 which is ~ected to create substantial shearing of the blood in the outer part of the rotor.
There thus remain a need for a VAD of simple design, small size, which enables a control mode and the potential for long term 5 durability, low heating and minimizes vibrations.
As well, there remains a need for such a VAD which restores hemodynamic function, avoids hemolysis and prevents clot formation, infection, and bleeding. A need for such a VAD of an implantable design also remains.
It would be beneficial to provide an axial flow blood pump which alleviates the limitations of the VAD of the prior art.
The present description refers to a number of documents, the colltelll of which is herein incorporated by reference.

The present invention seeks to meet these and other needs.
The present invention therefore seeks to provide an intra-ventricular assist device comprising an enclosed impeller axial flow 20 blood pump which restores hemodynamic function. In addition, the instant invention seeks to provide a VAD which avoids hemolysis. In addition, the VAD of the present invention seeks to prevent clot formation, infection and bleeding, and to minimize anti-coagulant requirements. Further, the present invention seeks to provide, in preferred embodiments, a VAD of 25 simple design, of a small size, and which enables a control mode and the potential for long durability (~ 2 years), low heating, and which minimizes vibration.

CA 02206644 1997-0~-30 The VAD of the present invention also seeks to provide a design thereof which minimizes the volumetric contribution from the hub, so as to maximize the volume of blood inside the pump.
The invention therefore features a ventricular assist device co",~,isi"g an axial flow blood pump provided with a stator and a rotor having enclosed impeller blades which intersect centrally to form a central hub at the axis of rotation of the rotor. The central hub is preferably designed to minimize its volumetric contribution to the rotor and hence to optimize the volume of blood pumped by the axial flow blood pump. The impeller blades are designed so as to minimize shearing forces, thereby minimizing hemolysis.
The invention further provides an implantable ventricular assist device comprising an axial flow blood pump provided with a stator and a rotor having enclosed impeller blades which intersect centrally to form a central hub at the axis of rotation of the rotor. The central hub is prererably designed to minimize its volumetric contribution to the rotor and hence to optimize the volume of blood pumped by the axial flow blood pump. The impeller blades are designed so as to minimize shearing forces, thereby minimizing hemolysis.
In accordance with one aspect of the invention, there is provided an axial flow blood pump which comprises a rotor having a cyli"dl ical member having an inner surface, the cylindrical member being provided with impeller extending inwardly from the inner surface of the rotor and intersecting at an axis of rotation of the rotor to form a central hub portion, with the central hub portion being rotatably mounted to a stator inside which the rotor rotates.

CA 02206644 1997-0~-30 In accordance with another aspect of the invention, there is provided a rotor having a cylindrical member having an inner surface, the cylindrical member being provided with impeller-extending inwardly from the inner surface of the rotor and intersecting at an axis of rotation of the rotor to form a central hub portion, with the central hub portion being rotalably mounted to a stator inside which the rotor rotates.
In accordance with a further aspect of the present invention, there is provided an intra-ventricular assist device comprising an axial flow blood pump which comprises a rotor having a cylindrical member having an inner surface, the cylindrical member being provided with impeller extending inwardly from the inner surface of the rotor and intersecting at an axis of rotation of the rotor to form a central hub portion, with the central hub portion being rotatably mounted to a stator inside which the rotor rotates.
In accordance with yet a further aspect of the present invention, there is provided a method for re~loring hemodynamic function and/or assist blood flow in a patient comprising the step of connecting a VAD in accordance with the present invention to the circulating system of the patient.
Other object, advantages, and features of the present invention will become more apparent upon reading of the following non-restrictive description of prefer,ed embodiment with reference to the accompanying drawings which are exemplary and should not be interpreted as limiting the scope of the present invention.

CA 02206644 1997-0~-30 DESCRIPTION OF THE PREFERRED EMBODIMENT
The originality of the proposed design is in part related to the axial flow blood pump itself, although the pump is only one element of the complete VAD system. In the following sections, we address 5 certain i",po,lant aspects concerning the pump design. In particular, the pump design takes into account anatomical and physiological considerations combined to mechanical, electrical and material requiremen~s. Then, following this discussion, the global characteristics of the VAD system are presented.
As alluded to above, it should be understood that the axial flow blood pump of the presenl invention should not be restricted to a use in an implantable VAD system. Indeed, it would be understood by the person of ordinary skill, that such a pump design which overcomes a number of drawbacks of the pump of the prior art, can be used as part of 15 an intra-corporal system such as an intra-ventricular VAD, or an extra-ventricular VAD (ie. abdomen), or alternatively as a para-corporal or extra~".oral VAD, often used in bridge to heart transplantation. It shall also be understood that the axial flow blood pump of the present invention can be used in temporary VAD's (ie. bridge to heart transplant) 20 or permanent VAD's, a non limiting example of which is the intra-ventricular VAD of the present invention.
What follows is an example of one type of permanent or temporary intra-ventricular VAD. For certainty, it shall be understood that the principle of the axial flow blood pump described herein is adaptable 25 to other types of VAD. Furthermore, it shall be understood that the intra-ventricular design desci ibed below is a preferred embodiment and hence CA 02206644 1997-0~-30 that a person of ordinary skill could adapt the present teachings to modify the VAD while remaining within the scope of the present invention.

ANATOMICAL PHYSIOLOGICAL AND SURGICAL CONSIDERATIONS
As previously discusserl bleeding is an important problem associated with patients who receive a VAD; in fact 30 % of patients suffer from this problem (Defraigne and Limet, 1 996b). The risk of infection is another quite i,npoi lant problem. Consequently, for medical and surgical considerations, a first objective was to have a completely intra-ventricular pump. Indeed this position eliminates the need for inflow and outflow grafts and their anastomoses which reduce the risk of bleeding and infection. This has also the obvious advantage of considerably reducing the implantation time. Figure 9 illustrates the proposed position of the pump in the left ventricle.
We have opted to design a pump that could fit in small adults and in teens. Thus we limit the size of the pump in regard to the ventricular dimension of people with a BSA (Body Surface Area) of 1.5 m2. Since the physical size and shape of the pump are greatly influenced by the desir~d location of the pump, a good description of the ventricle anatomy is required. Feigenbaum (1994) presents several dimensions of the heart normalized by the BSA (Body Surface Area).
These analorr, --' dimensions have been statistically determined and are known to represent 95 % of the population. Thus we used the ventricular dimension of people with a BSA of 1.5 m2.
It will be understood that the physical size and shape of the pump could also be adapted to meet the anatomical dimensions of individual falling outside this 95% of the population. Similarly, the size CA 02206644 1997-0~-30 and shape could be adapted to specific individuals and heart conditions as known to the person of ordinary skill.
The internal diameter of the left ventricle ranges from 37 to 46 mm in diastole and between 22 to 31 mm in systole. This diameter 5 is determined in the middle of the ventricular length (segment AB in Figure 9). The diameter near the apex at the first third of the ventricle length is about 1.5 cm (segment CD). The internal length of the ventricle from the apex to mitral valve ranges from 55 to 70 mm. Finally, the other important parameter is the surface of the aortic valve opening which 10 ranges from 2.5 to 4 cm2. In certain situations, the valves may be the site of an obstruction (stenosis). It is interesting to note that an opening of only 1.1 cm2 is considered a light slenosis. As it will be discussed in more details later, these anatomic figures were very important for the design of pump and VAD's of the present invention.
In a surgical point of view, the favored insertion procedure is to use the same approach as with valve replacement. An incision would be made at the root of the aorta and the pump would be p~ssed though the aortic valve and then into the left ventricle. The pump would be pushed until its base reaches the myocardium at the apex. In 20 order to prevent its motion, the pump could be fixed. As will be seen below, in a prefer,ed embodiment, the pump is fixed at the apex. In a prefer,ed embodiment, an oufflow cannula could then pass through the aortic valves to further reduce hemolysis.
As previously mentioned, one of the first role of the 25 pump is to restore hemodynamic function in patients with cardiac failure.
Depending on the severity of the failure and the BSA, the pump should work at flow rates between 1 to 10 L/min against a pressure as high as CA 02206644 1997-0~-30 200 mmHg and preferably at a flow rate between 3 to 5 Umin against a pressure of 100 mmHg.
Another impo, lanl co"sideralion for blood pump design is the hemolysis rate. Hemolysis is the tearing of red blood cells which 5 empties the contenl of the cells in the blood stream resulting in free hemoglobin; the normal level of plasma free hemoglobin is around 10 mg/dL. A normalized index of hemolysis (NIH) of 0.005 g/100 L and lower is considered to be almost athromatic for red blood cells. A NIH of about 0.05 g/100 L could be tolerated. A NIH of between 0.05 g/100 L to 0.005 10 g/100 L is envisaged for the VAD of the present invention. Preferably, the NIH is in the 0.005 g/100 L range. The platelets are another important blood elements, their activation by high hydromechanical forces should be avoided in order to prevent clot formation.

Mechanical aspect This section is divided into two parts. The first part desc, ibes the extemal shape and sizes of a preferred embodiment of the pump and the second describes the internal conception thereof.
Extemal design requiren~e..ls The external design of the VAD depends on the anatomic dimensions of the left ventricle. Figure 10 illustrates the external shape of the pump and the critical geometric parameters.
Inflow cage - From Figure 9 it is seen that with the proposed location, the base of the pump would sit at the bottom of the heart. Consequently, CA 02206644 1997-0~-30 in order to prevent the ventricle walls from obstructing blood intake, a cage is dlla~,ed to the extremity of the pump. The diameter of the inflow cage is set to 15 mm which is smaller than the segment CD (Figure 9) and should limit the pressure on the ventricle wall near the apex.

Oufflow cannula - At the other end of the pump, the outlet area is reduced so that the surface of the outflow cannula minimally obstructs the aortic valves. BecA!~se we are interested in assisted circulation, blood flow contribution from the natural contraction should be maintained. As 10 previously "~er,lioned the aortic valve openings ranges from 2.5 to 4 cm2 and a valve opening reduced to 1.1 cm2 mildly affects normal blood flow.
Consideri~ ,g these figures, an obstruction smaller than 1.4 cm2 (2.5 cm2 1.1 cm2) should minimally disturb the valve function. In a preferred embodiment, an oufflow cannula of 39 F, that is 1.3 cm2 was chosen.
15 Moreover since the cannula is not an obstruction, it should be even less disturbing for blood flow.
At the distal end of the cannula, a diffuser may be present. This diffuser has the function of reducing the shear stress on blood cells. Without a diffuser, the blood velocity ejected by the pump is 20 greater than the one ejected by the heart. The difference in velocity between these two flows will result in shear stress proportional to this difference. Since, the velocity is inversely proportional to the cross-section area, in order to reduce the relative speed, we can increase the area occupied by the pump cannula orifice (reduce its speed) which 25 decrease the area occupied by the blood flow (increase its speed). This is exactly the role of the proposed diffuser. The opening angle and the CA 02206644 1997-0~-30 length of the diffuser will be adjust upon me~;han.c ~I characteristics of the pump to minimize the shear stress.

Housing - The dia",e(er of the pump is a co",promise between pumping 5 requirements and minimal interference with heart contraction. In a prefe"ed embodiment the maximum allowable diameter is about 22 mm which is the ventricle diameter in systole. This dimension is reasonable since people with heart failure have dilated ventricles.
The maximum length of the pump as illustrated in 10 Figure 10 is set in regard of the average distance between the apex and the mitral valve. In a prefer,ed embodiment the length of the pump is chosen as about 55 mm. As shown a reduction of the pump diameter toward the outflow increases the mitral valve clearance in order to minimize inte, rerence with the mitral valve function.
Since in this particular embodiment the pump will be completely inside the ventricle blood will circulate around its casing. In consequence external surfaces should be smooth and not have abrupt changes to minimize vortices turbulence and recirculation zones which may be at the origin of clot formation.
Fixation mechanism - At the pump inflow a fixation mechanism is propose~l This mechanism would pass through the myocardium and on the epicardium; a sewing ring would be provided to firmly fix the pump in place.

CA 02206644 1997-0~-30 Electrical supply - The required electrical supply for the motor operation would be carried by a wire that could, for example, pass by the fixation mechanism to the controller and to the energy source.

5 Intemal design characlerislics In the design of the internal components, the general principles of axial flow blood pump have been retained. At the inflow, there is a stator (a slalionary part) that induces a rotation to the flow. This initial rotation minimizes abrupt changes in the blood flow path when it 10 reaches the impeller (rotor). At the oufflow, another stator is used to ll dl ,srul ", the rolatio, ~al motion into a translational motion; in other words, the stator acts as a flow straightener.
As previously mentioned, the pump design should minimize shearing stress in order to minimize hemolysis. In that context, 15 reduction of the rotational speed (RPM) would obviously contribute to reduce hemolysis. This implies that the pumping volume per revolution should be increased to keep the required blood flow. The pumping volume corresponds to the volume of blood contained in the pump rotor zone. A way to increase the pumping volume is to minimize the volume 20 of the central hub of the pump rotor. As evidenced herein, this aspect is an important aspect of the present invention.
Indeed, in current wet motor axial flow blood pump there is two ways to fix the permanent magnets. The first is to put them in the central hub and the second is to insert them in the blades. Both methods 25 require important compromises. With the first method a large hub is required to locate the permanent magnets close to the motor windings for electrical coupling reason. In cor,l,dsl, a small hub is required to increase CA 02206644 1997-0~-30 the pumping volume. The second yields a compromise since the geometry of the blades must be curved for pumping efficiency. As a cons~uence, some of the er,~bedded magnets get farther away from the windings as the blade is curving (Figure 7).
The approach proposed herein eliminates these co,nproi,lises allowing an optimal electro-magnetic coupling and an optimal pumping volume at the same time. We propose an Enclosed-l,llpell¢rAxial Flowconfiguration: the blades are enclosed in a casing so that they are locked with one another. This allows a reduction of the central hub for optimal pumping volume while providing the required structural supports for the blades (Figure 12). In addition, the permanent magnets can then be inserted in the blades casing, providing optimal coupling.
At the blades intersection a shaft protrudes on each extremity of the impeller. This shafts will serve to support the rotor and act as the bearing system. In one embodiment, the two extremities of the shaft are insel led into two rings located on the stators. The small central hub should also allow good beal ing washout. Indeed, the bearing system is a critical part of the pump, since the bearings are the only mechanical parts subject to wear, it is the part which is expected to be mostly responsible for the life of the pump.
The number of blades and their angulation will initially be determined on the basis of hydrodynamic relations and then optimized with CFD simulations. The CFD approach is r~iscussed in more details below.
Figure 13 illustrates that the blades casing is contiguous with the inside walls of the pump casing. In other words, the internal CA 02206644 1997-0~-30 diameter of the blades casing is the same as the internal diameter of the stators, which minimizes abrupt changes responsible for flow perturbations.
In the proposed configuration, the closed rotor (blades 5 and blade casing) is insei led in the pump casing, this could be viewed as a double casing configuration. The gap that separates these two elements should be as small as possible to minimize the penetration of blood elements in this area but not too small to prevent shocks due to rotor vibrations.
In order to minimize blood flow in this junction, an hydraulic seal may be used. In an embodiment, the seal can be made by a set of tiny grooves at the surface of the blade casing cylinder and at the intemal surface of the pump casing. Furthermore, the angulation on each side of the cylinder (Figure 13) is oriented in order to minimize inflow of 15 particles.

Electrical AsPects The mechanical pumping will be actuated by means of a brushless DC motor. The brushless configuration presents the 20 advantage of minimal wear. Two other interesting characteristics of brushless DC motors are high rotational speed and high torque. The motor is cori,posed of two elemen~s: permanent magnets in the rotor and windings in the stator.
As discussed previously, the configuration where the 25 blades are enclosed in a tube to increase the pumping volume has also an advantage related to the electrical coupling of the motor. In the proposed design, the permanent magnets are embedded in the blades CA 02206644 1997-0~-30 casing of the rotor and the windings are embedded in the enlarged part of the pump casing over the rotor. The presence of the permanent ",ay"ets in the blades casing very close to the winding provides an optimal ele~,omaynelic coupling and obviously minimizes electrical loss.
5 When co",pared to the configurations with permanent magnets enclosed in the blades, the proposed configuration allows a number of magnet independent of the number of blades.

Materials selection The choice of materials for an implantable system is crucial and several properties of the materials available should be considered: strength, durability hardness, elasticity, wear resistance, surface finish and biocompatibility. Biocompatibility is very important to minimize irritation, rejection and thrombogenesis. The interaction 15 between the surface of the material and the biological tissues is very complex. In several cases, treatments of the surface with human protein or with certain drugs like heparin may considerably increase the biocGIllpdlibility and minimize thrombus formation.
The following table SUm"~a, i~es both the properties and 20 a selection of exemplary materials required for the different elements of the pump.

CA 02206644 1997-0~-30 Table 3: Examples of materials and their properties which can be selected for the different elements of the pump.

Components Properties Materials Cage - Hemocompatibility - Titanium alloys (6AI - 4V) - Anti-inflammatory - Rigidity Casing - Hemocompatibility - Titanium alloys - Anti-inflammatory - Rigidity - Weight Outflow cannula - Flexibility - Dacron grafVGore-Tex graft - Smooth surface Impeller and flow stator - Hemocompatibility - Titanium alloys - Rigidity - Weight Shaft and bearings - Wear resistance - Sapphire and zirconia - Hardness - Rigidity Surface l,eal",e.. t - Hemocompatibility - Heparin GLOBAL SYSTEM
Figure 14 illustrates the configuration of an embodiment of a totally i" ,planlable axial flow system. The system is composed of four main parts: 1- an axial flow blood pump, 2- an internal controller, 3- two energy sources and 4- a Transcutaneous Energy and Information 20 Transmission (TEIT) system.
The TEIT system works on the same principles as a l,;ansror",er. It is composed of two electrical coils, one external and one internal. A RF (Radio Frequency) signal is applied to the external coil CA 02206644 1997-0~-30 (transmission coil) and the signal is transmitted to the inler"al coil (reception coil). The transmitted energy will be used to supply all the co,nponenls of the VAD system. The same TEIT system can also be used for transferring operational data and progra",mable para",eters.
The VAD system includes three external components that could be i"co"Joialed to a belt or suspenders: The first element is a source of energy in the form of a rechargeable battery for good mobility (the source could also be fixed when the patient is stationary for example during sleeping time). The second element is a transmission circuit that converts the DC voltage from the energy source into an RF signal. Finally a third element the l,dns,nission coil would transfer the RF signal to the reception coil.
The VAD system also requires the implantation of four inle",al cc"~po,1e"ls. Thefirst component is the pump itself. The second component is a reception coil placed in a sub-cutaneous position as described above. The third component is a rechargeable battery that should provide a complete autonomy of more than 30 min. As a consequence when the patient removes his belt or suspenders the inlei"al battery will provide the required power. The forth element is the controller. The controller is a very important part of the system and achieves several functions. Its main function is to determine the required blood flow rate that must be delivered by the pump considering the BSA
and the level of activity of the patient (estimated by the heart rate).
Depending of the measurements the controller will adjust the pump speed in consequence. Most of the other functions of the pump are related to power manage" ,enl. The controller converts the RF signal from the inte" ,al coil in DC. This DC is used by the controller to supply power CA 02206644 1997-0~-30 to the pump and to recharge the internal battery when required. The controller also monitors the pump energy consumption. Interestingly, perturbation in energy consumption may serve to diagnose mechanical problems. For e~a,nplc, a slow but constant increase may reveal bearing 5 wear or U llolllbus formation interfering with the pump operation. Another possi'~ilily is a sudden inuease in energy consumption which may reveal a significant obstruction of the pump. The last function of the controller is to communicate inrorl"alion. In that context it is used to receive external programmable parameters and to transfer monitoring data to a 10 computer in order to be processed.
BecAuse the controller is the central element of the VAD
system, its reliability is imperative. To minimize the risk of hazards and malfunctions, several strategies must be combined. First, the redundancy of critical electronic components is recommended, a watch dog circuit 15 that trap erratic behaviors of the controller is proposed. Finally, we shouldcouple to the watch dog circuit a backup circuit that would provide a fix blood flow in case of major problems with the controller.

DESIGN EVALUATION
The evaluation process is a crucial aspect of the present project and will directly affect the design of the pump. The process and the tools used are determinant on the validation of the pump design.
25 Once the initial design is made, the evaluation process will be initiated and it should lead us toward an optimal design according to the diagram shown in Figure 15. The three evaluation phases are Computational Fluid CA 02206644 1997-0~-30 Dynamic (CFD), in-vitro experiments and in-vivo experiments. The optimal design shall be reached by an iterative process between the evaluation phases and the design phase.
At each evaluation phase, a set of observations and 5 measurements will be made. After each phase, based on the observations done, we will choose either to continue with the next evaluation process if all requirements are met, or to proceed with design modifications in order to improve the pump characteristics. All three phases are important and are discussed below.
CFD: Numerical evaluation The use of advanced numerical techniques for characterizing the performance of a blood pump is a cost-effective and viable alternative to experimental evaluation. A numerical model of the 15 pump geometry can be created using computer-aided design software.
With the help of computational fluid dynamics (CFD), the entire flow field in the pump can be simulated and analysed under physiologically relevant conditions. With the use of this methodology, the flow region is divided (discritized) into small computation cells with the help of a grid 20 (mesh). The equations of mass, ~lo~enlum and energy conservation that govem fluid motion are expressed in an algebraic form. These take into account the physical conditions specifled at the flow domain boundaries.
The solution of these equations is obtained numerically in each cell.
Based on the results of the analysis, different design modifications, such 25 as changes in the impeller configurations, can be implemented on the numerical prototype. The effect of these modifications on the pump can be studied further with the CFD model. Subsequently, such parametric CA 02206644 1997-0~-30 analyses can be pe, ror"~ed on multiple designs to evaluate the merits of each design.

Blood flow dynamics Considering the physical size of the pump and the blood flow rate through the pump, blood can be modelled as a Newtonian fluid.
Indeed, the dimension of the red blood cells can be neglected, blood can be considered as a continuous fluid and the velocity of the blood yields relatively large Reynolds numbers. Within the pump the flow is clearly unsteady. However, the complexity of the flow can be broken down into a hierarchy of approximations:
A through flow analysis with a distributed loss model.
A passage-averaged analysis coupling 3D stator-rotor interaction.
An aperiodic transient passage-averaged analysis.
The first two of these approximations are steady state approximations of the Navier-Stokes equations:

-- + (u . ~u~= -Vp + 1 ~2u~
~ R
V. U =O

where is the velocity vector, u the pressure and R is the Reynolds number which reflects the relative importance of convective forces to viscous forces. The equation R=UL/v defines the Reynolds number with U the velocity magnitude of the incoming blood flow, L the vessel diameter and n the cinematic viscosity. The flow is assumed CA 02206644 1997-0~-30 incompressible and the main physical constants are the blood density Pb=1 .17 g/cm3 and the blood static viscosity ,Ub=3.5 X 1 o-2 poise.

Numerical solution methodology Two commercial codes can be used: FIDAP (finite element) and STAR-CD (control volume). The latter is interesting for its moving grid capability used in detailed rotor-stator computations. Finally, an academic code EF (finite element) will also be used because of its unstructured grid car~hility which allows grid adaptivity and optimal error control. All these codes are using the SIMPLER algorithm with a transient implicit or semi-implicit formulations. Furthermore, the momentum equations are segregated to maximize the number of grid points. A
proven particle tracking model developed at École Polytechnique ~ 10n~réal, Québec) will be used to study the aggregation tendency of blood particles in a particular region of the flow domain and to estimate their residence time in areas of high fluid shear stress. These data can be used to predict hemolysis and thrombus formation in the pump. The CFD simulations could also be combined to hemolysis model such as Pinotti and Rosa (1995), to have an idea of the hemolysis rate of a particular design.
The blade and the pump geometry will be modelized preferably with a parametric model such as ProEngineer or CATIA. These modelers generate an IGES file which can be read into any bf the above codes. This geometry program will allow to investigate the influence of the blade angulation on the level of shearing stresses. The pump geometry will also be investigated for any turbulence and cavitation a CA 02206644 1997-0~-30 phenomena also responsible for red blood cell damage and increase of hemolysis rate.

In-vitro ex~eri",e..ts The objectives of the in-vitro experiments (P.M. Galletti, 1995) are to: 1- Observe oper~lion and assess pe, rO" ,~ance of the device under controlled conditions; 2- Define performance in quantitative terms over a wide range of input conditions; and 3- Assess device reliability and durability.
To acl ,.eve these objectives, three in-vitro experiments are planned to evaluate the pump cha,~~teristics, they are hydrodynamic pe,ruilllancesl hemolysis and wear test. These three characteristics will be evaluated by the experiments described below.

Hydrodynamic characteristics First the hydrodynamic characteristics of the pump will be evaluated; that is the pressure vs flow characteristics for different pump .speed.s This will be done with a simple hydraulic loop and a blood analogue. The blood analogue is composed of water (63 %) and glycerin (37 %) which has a density and a viscosity close to those of blood. This simple hydraulic loop will be composed of a fluid reservoir, a cannula connecting the reservoir to the pump inflow, the pump and an oufflow cannula connected to the fluid reservoir. A variable flow restrictor in the outflow cannula will allow to vary the hydrodynamic resistance of the circuit. Pressure transducer at pump inflow and oufflow will allow to measure pressure dirrere"ce and a flow transducer to measure pump CA 02206644 1997-0~-30 flow. This e~ri" ,enl will also serve to estimate energy consumption and heat generation.

Hemolysis rate To assess hemolysis, the circuit described above will be used with fresh bovine blood as circulating fluid. Blood at 37 ~C will be pumped at a rate of 5 L/min against a pressure of 100 mm Hg. A
thermostatic water bath set at 37 ~C will maintain blood temperature. In order to compare hemolysis level to other pump, the NIH (Normalized Index of Hemolysis) will be computed.

NIH = ~IHb x V x (1 - Ht) x 100 Qx T

The NIH (9/100 L) is defined as the amount of hemoglobin in gram released per 100 L of blood pumped. The parameters are: T, the pumping time in minute, Q, the flow rate in L/min, ~Hb, the plasma free hemoglobin increase (g/L) during the pumping time, Ht, the hematocit and V, the volume of blood contained in the circuit.
Prior to begin the NIH evaluation with the axial flow blood pump designed, the hydraulic circuit will be used with the Bio-Medicus pump, a pump known to produce a low hemolysis rate. This will allow to isolate the hemolysis level caused by the hydraulic circuit itself.
At the end of this experiment the pump will be ~lis~-ssembled and the presence of blood clots between the blades casing and the pump casing will be quantified in order to evaluate the efficiency of the hydraulic seals.

CA 02206644 1997-0~-30 Wear characteri~lics Accelerated wear e~erilllents are standard procedures to evaluate mechanical components and are particularly critical for implants with moving parts subjected to repeated stress (P.M. Galletti, 1995). The goal of this in-vitro experiment is to estimate the wear chara~1eri~lics of the pump. A blood analogue solution will be pumped at an higher speed to accelerate the wearing process. Power consumption will be monitored as a first estimate of the wear. Disassembling of the pump will allow to assess more accurately the wear, in particular at bearing location.
In-vitro experiments are required to validate CFD
results, assess pump pei ro, mance and have knowledge of the interaction mechanisms between the pump and blood elements. However, since it is difficult to maintain blood in a normal physiologic state for a long time in-vitro conditions, in-vivo experiments are required to evaluate the impact of the pump on physiological tissues and on thrombus formation over a long period of time.

In-vivo exPeriments The objectives of the in-vivo experiments are also multiple that is: 1- Evaluate the biocompatibility of the device and its impact on organ function; 2- Evaluate the effectiveness of the device to replace a physiologic function and related adverse effects; and 3-Develop the surgical procedure and post-surgical management of the device receiver. These objectives could be achieved with animal studies using short term and long term experiments.

CA 02206644 1997-0~-30 Short terrn animal studies The goals of these experiments are to evaluate short term impact of the pump implantation in a living body. The experiments will be pe, ro""ed on 5 healthy dogs wci~hling at least 30 kg. The animals will be instrumented and the experiments will last 30 days. Since the animals will be instrumented, it is more practical to control the pump externally through a percutaneous wire. The parameters to be investigated are:
The surgical procedures for the implantation of the pump.
The pump cha,a~leristics (pulsatile pressure flow-rate versus pump speed).
The post-surgical management of the receiver animal (blood sample analysis for drug therapy and anti-coagulant will be studied).
The control algorithm under physiologic conditions. The algorill,i" should adjust the pump speed in order to produce the proper blood flow.
The hemolysis rate with the plasma free hemoglobin.
Organ dysfunction (e.g. renal function).
Upon completion of the experiment complete autopsy of the animal for organ dysfunction or thromboembolism.
Clot for"lalion site by dis~ssembling the pump.
Necrosis or inflammation of tissue in contact with the pump.

Long term animal studies The goals of these experiments are to evaluate the potential of the pump as treatment of heart failure and the biocompatibility of the pump over a long period of time. Experiments will be performed on 5 dogs (30 kg and more) to which heart failure will have been induced.

CA 02206644 1997-0~-30 The purpose is to ~ssess the longevity of animals with circulatory assist device and to monitor their health conditions. Experiments will be terminated only if a pump defect occurs or if multiple medical complications occur. The parameters to be investigated are:
Evaluation and treatment of right heart failure due to left ventricular assist support.
Pump durability in in-vivo conditions.
Long term treatment of heart failure with the axial flow blood pump designed.
Biocompatibility of the pump material.
Upon termination of the experiments, autopsy will be performed and pump examined.

CONCLUSIONS
Ventricular assist devices are now being used world wide and their utilization is becoming more and more accepted as a solution to treat end stage heart failure. It is being accepted that VAD
extend life extent and provide good increase in quality of life of patients.
A pool, made with patients who received VAD, concerning their quality of life rcve-'ed that these patients would have preferred a heart transplant but prefer their situation than having to be on dialyses.
It is also now being accepted that VAD is becoming a cost effective solution considering the fact that patients are discharged from the hospital more rapidly and may return to normal life occupations.
In United-States, several insurance companies are now reimbursing the implantation of such device.
VAD has shown good results in the past as a bridge to heart transplant but the bridge solution may only postpone the problem CA 02206644 1997-0~-30 of the lack of heart donors. Consequently, the new goal of VAD design is to develop permanent implants, which could also be a less costly solution.
Although VAD appears to be a good appr~acl), there are 5 still several remaining problems ~ssoci-te~ with actual VAD. Pulsatile VAD which are now currently available are related to technologies developed in the 70'. They produce high hemolysis rate, have high risk of infection and high risk of bleeding. They are also very large, heavy (about 2 pounds), mechanically complex and costly.
Recently, developed axial flow blood pumps alleviate some of the limitations related to size and weight, and infection but problems such as hemolysis and bleeding are still remaining. It is ex~ ~e.:led that the Enclosed-lmpeller Axial Flow blood Pump geometry of the present invention will contribute to go one step further toward an 15 optimal solution. This could be achieved with the important conl,ibution of Computational Fluid D~,la(r,.cs. CFD simulations will allow to evaluate performances of the pump even before the pump is manufactured.
Finally the proposed axial flow blood pump should provide an excellent bridge to heart transplant and aims at long term 20 implant. Indeed, the proposed technology should answer most of the remaining limitations with actual axial flow blood pump and especially hemolysis and bleeding.
Various modifications of the invention in addition to those shown and described herein will become apparent to those of 25 ordinary skill from the foregoing description and accompanying drawings, such modifications are intended to fall within the scope of the appended claims.

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Claims (6)

1. An axial flow blood pump comprising:
- a rotor having a cylindrical member defining an inner surface, said cylindrical member being provided with impellers extending inwardly from said inner surface and intersecting at an axis of rotation of said rotor to form a central hub portion;
- a stator enclosing said rotor; and - means for rotatably mounting said central hub portion to the stator.
2. A ventricular assist device (VAD) comprising said axial flow blood pump of claim 1.
3. The VAD of claim 2, wherein said VAD is implantable inside the body of a patient in need thereof.
4. The VAD of claim 3, wherein said VAD is intra-ventricular.
5. A method for restoring hemodynamic function in a patient in need thereof, comprising the step of connecting said axial flow blood pump of claim 1 to the circulatory system of said patient so as to enable pumping of blood in said patient's bloodstream.
6. A method for assisting blood flow in a patient in need thereof, comprising the step of connecting said axial flow blood pump of claim 1 to the circulatory system of said patient so as to enable pumping of blood in said patient's bloodstream.
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