Heart pump outflow conduit

Abstract

An outflow conduit for a heart pump, the outflow conduit including an inlet configured to be attached to a pump outlet of the heart pump to receive blood therefrom, an outlet configured to be connected to an arterial vessel, wherein the outlet includes a flared portion to minimise flow restrictions and a sidewall defining an internal cavity extending from the inlet to the outlet, the cavity being configured to transport blood from the heart pump to the arterial vessel.

Description

HEART PUMP OUTFLOW CONDUITBackground of the Invention

[0001] The present invention relates to an outflow conduit for a heart pump, and in one particular example, an outflow conduit for connecting an outlet of a heart pump, such as a ventricular assist device, to an arterial vessel.Description of the Prior Art

[0002] The reference in this specification to any prior publication (or information derived from it), or to any matter which is known, is not, and should not be taken as an acknowledgment or admission or any form of suggestion that the prior publication (or information derived from it) or known matter forms part of the common general knowledge in the field of endeavour to which this specification relates.

[0003] Patients with impaired ventricular function or severe heart dysfunctions require mechanical assistant in blood circulatory. Mechanical circulatory support is an effective means to restore blood flow to the circulatory system in the case of a failing heart. This support can be in the form of a ventricular assistance device (VAD), whereby the device assist the function of one or more of the failing ventricles, or complete replacement of the ventricles with a total artificial heart (TAH).

[0004] For example, left ventricular assist devices are pumps that attach to the cardiac apex and deliver blood via pump conduit to the ascending aorta. They are used in patients with severe heart failure and completely support the circulation in -70% of cases, resulting in complete closure of the aortic valve.

[0005] Currently, most heart pumps are designed to operate with an optimum operating efficiency at a normal flow rate for a subject, which typically corresponds to a flow rate of about five to six litres per minute. However, such configurations often have a steep pressureflow curve, which in turn makes the pumps sensitive to pressure changes. This lowers pump efficiency and means that a large increase in pump speed may be required in order to counteract pressure changes.

[0006] To date, the majority of work to address this has focused on improving pump hydrodynamic and mechanical efficiency, for example by improving the mechanical or electrical design of the pump, so that the pump can handle a range of flow conditions while operating in a safe pressure condition.

[0007] In order to attach a heart pump to the circulatory system a range of different conduit designs have been proposed.

[0008] An example of a conduit is disclosed in US 9585991. The document discloses an apparatus includes a tubular member defining a lumen and a channel, and a support member. The tubular member has a connection portion configured to be coupled to an organ wall. An outer edge of the connection portion is configured to contact a first portion of an inner surface of the wall when the connection portion is in an expanded configuration such that an interior volume of the organ is in fluid communication with the lumen. The support member is movably disposed within the channel and is configured to minimize movement of the wall relative to the tubular member. An end portion of the support member is disposed within the channel when the support member is in a first configuration. The end portion of the support member configured to contact a second portion of the inner surface of the wall when the support member is in a second configuration.

[0009] A further example is disclosed in US 2012 / 0296152, which relates to methods and materials for providing blood flow for a blood pump recipient. For example, cannulae that can be connected to the circulatory system of a mammal and can be used in conjunction with a blood pump (e.g., an assist device) are provided.

[0010] A further example is disclosed in CN 202920776. The document relates to an inflow pipeline of an axial flow type left heart auxiliary pump, which comprises a bell-mouth- shaped inlet turned over externally, a suture ring, a pipe body and an outlet. Aiming at the requirement of implanting operation of the axial flow type left heart auxiliary pump, the problems of the axial flow type left heart auxiliary pump caused by cannulation of venae cava during the operation are solved. The inlet of the inflow pipeline is in a bell-mouth-shaped flaring structure, and the inlet after implantation is attached to a ventricular wall, so that the inlet can be efficiently prevented from absorbing the ventricular wall, the damage of blood inside a ventricle and the pipeline is reduced, and the risks that the mouth of the pipe is pulled out andthe blood at the suturing part is oozed can be efficiently reduced through the suture ring on the outer wall of the pipeline. The inflow pipeline can cause the axial flow type left heart auxiliary pump to provide better function of pumping blood.

[0011] However, none of these conduits are specifically designed with regard to pump efficiency.Summary of the Present Invention

[0012] In one broad form an aspect of the present invention seeks to provide an outflow conduit for a heart pump, the outflow conduit including: an inlet configured to be attached to a pump outlet of the heart pump to receive blood therefrom; an outlet configured to be connected to an arterial vessel, wherein the outlet includes a flared portion to minimise flow restrictions; and a sidewall defining an internal cavity extending from the inlet to the outlet, the cavity being configured to transport blood from the heart pump to the arterial vessel.

[0013] In one embodiment the outlet is configured to disperse blood flow into the arterial vessel.

[0014] In one embodiment the flared portion increases in internal diameter towards the outlet.

[0015] In one embodiment the flared portion includes a starting diameter and an ending diameter, wherein the ending diameter is larger than the starting diameter.

[0016] In one embodiment the ending diameter is at least one of: about 1.2 to 1.6 times the starting diameter; about 12 mm to 16 mm; and about 14 mm.

[0017] In one embodiment the starting diameter is at least one of: approximately equal to an internal diameter of the sidewall; and, about 8 mm to 12 mm; about 10 mm.

[0018] In one embodiment the flared portion is at least one of: about 10% to 30% of an outflow conduit length; about 25 mm to 35 mm in length; and about 30 mm in length.

[0019] In one embodiment the outlet is configured to be connected to an arch of the arterial vessel.

[0020] In one embodiment the outlet is at least one of: thicker than the sidewall; and stiffer than the sidewall.

[0021] In one embodiment the outlet includes at least one surface structure to maintain a shape of the outlet.

[0022] In one embodiment the outlet includes a rim.

[0023] The flared portion is defined by a flared profile of the rim.

[0024] In one embodiment at least part of the sidewall is coated.

[0025] In one embodiment coating is at least one of: a low friction coating; and an antibiofouling coating.

[0026] In one embodiment the conduit is made of an at least partially compliant material.

[0027] In one embodiment the conduit is made from at least one of: nanofibers; electrospun nanofibers; silk; electrospun silk; and electrospun silk with a nanocoating.

[0028] In one embodiment the conduit includes a drug eluting conduit.

[0029] In one embodiment the drug inhibits at least one of: inflammatory response; and immune system responses.

[0030] It will be appreciated that the broad forms of the invention and their respective features can be used in conjunction and / or independently, and reference to separate broad forms is not intended to be limiting. Furthermore, it will be appreciated that features of the method can be performed using the system or apparatus and that features of the system or apparatus can be implemented using the method.Brief Description of the Drawings

[0031] Various examples and embodiments of the present invention will now be described with reference to the accompanying drawings, in which: -

[0032] Figure 1 is a schematic diagram of an example of a heart pump;

[0033] Figure 2 is a schematic diagram of an example of an outflow conduit for use with a heart pump;

[0034] Figure 3 is a schematic diagram of an outlet of an outflow conduit;

[0035] Figure 4 is a schematic diagram of an outflow conduit being attached to an arterial vessel;

[0036] Figures 5A and 5B are contrast-enhanced axial plane computerized tomography demonstrating subtle stenosis in an outflow graft-aorta anastomosis;

[0037] Figures 5C and 5D are contrast-enhanced sagittal plane computerized tomography demonstrating subtle stenosis in the outflow graft-aorta anastomosis of Figures 5A and 5B;

[0038] Figure 6 is a graph illustrating examples of pressure loss within different outflow conduits; and,

[0039] Figures 7 A to 7C are images of example of flow streamlines for different outflow conduits.Detailed Description of the Preferred Embodiments

[0040] An example of a heart pump will now be described with reference to Figure 1.

[0041] In this example, there is provided a heart pump 100 that is connected to the heart 150 of a subject. The heart pump 100 is coupled via respective inflow conduit 110 and outflow conduit 120 to the left ventricle 151 and aorta 152, and is therefore functioning as a left ventricular assist device (LVAD), although this is not essential and similar techniques to those described can also be applied to right ventricular assist devices (RVADs) coupled to the right ventricle and pulmonary artery.

[0042] In this example, the heart pump is a continuous flow VAD (cfVAD) in which an impeller is continuously rotated within a cavity, to thereby pump blood from the ventricle into the aorta. The heart pump 100 can be a standard VAD known in the art, such as a HeartWare HVAD, Ventracor Ventrassist, or the like, and this will not therefore be described in further detail.

[0043] A controller 160 may also be coupled to the heart pump 100 that operates to control the heart pump 100 and in particular control rotation of the impeller and optionally monitor operating characteristics of the heart pump 100.

[0044] An example of an outflow conduit for use with a heart pump will now be described with reference to Figure 2.

[0045] In this example, there is provided an outflow conduit 220 for a heart pump 200. The outflow conduit 220 includes an inlet 221, an outlet 222 and a sidewall 223 defining an internal cavity extending from the inlet 221 to the outlet 222. The inlet 221 is attached to a pump outlet of the heart pump 200 to receive blood therefrom, whilst the outlet 222 is connected to an arterial vessel 252, such that the cavity transports blood from the heart pump 200 to the arterial vessel 252.

[0046] In this example, the outlet 222 includes a flared portion 222a. The flared portion 222a is configured to try and minimise flow restrictions in the outlet conduit, which can in turn lead to afterload pressures in the outlet conduit, which in turn has an impact on pump efficiency and efficacy.

[0047] In this regard, it has been identified that stenosis in the conduit, particularly in the outlet, can arise for a number of reasons. For example, extrinsic compression, kinking, or twisting may arise as the conduit is positioned within the patient, whilst anastomotic narrowing can arise when a surgeon attaches the conduit to the arterial vessel. Additionally, these issues are further exacerbated over time by fibrosis during the healing process, which can lead to tissue ingrowth where the conduit and arterial vessel meet, as well as thrombus formation, or the like, which may put further restrictions on the blood flow within the conduit, and particularly at the outlet.

[0048] It has been found that in traditional arrangements, narrowing in the outlet increases pressure in the conduit and leads to a higher pump afterload, which can in turn significantly reduce pump effectiveness. This leads to a reduction in pump effectiveness, which increases over time as ingrowth continues. Whilst this can be counteracted by operating the pump at a higher rate, this in turn reduces battery life and increases wear, reducing pump life.Additionally, this can lead to increased stress on the blood, which in turn may lead to adverse consequences, such as hemolysis.

[0049] Furthermore, narrowing within the outlet can give rise to pulsatile effects, and specifically pulsatile inertance, which is an additional impediment to pulsatile flow. In this regard, studies such as “Pulsatile Conduit Pressure Gradients in the HeartWare HVAD” by Pankaj Jain, Sajad Shehab, Michael Stevens, Peter Macdonald, Paul Jansz, And Christopher Hayward, ASAIO Journal 2019, July 2019 - Volume 65 - Issue 5 - p 489-494, have shown mean pressure gradients across the current conduits of approximately 20mmHg during normal pump function, with gradients increasing up to 40mmHg at peak flow (8L / min). This not only further decreases the average flow support provided by the pump but also further limits increase pump flows which are typically required during exercise, or the like.

[0050] Finally, narrowing within the conduit outlet can lead to a jetting effect, in which blood is effectively forced out of the conduit in a jet stream. This can impact on the blood flow within the vessel, specifically leading to enhanced shear and imbalanced flow in aortic arch vessels, which in turn leads to a high degree of sensitivity to graft insertion angle. This can also directly impinge on the blood vessel, leading to shear stress that damages endothelial cells and leads to phlebitis (inflammation). To counteract this, an insertion angle of the conduit into the arterial vessel typically needs to be carefully controlled so that any jet is directed into the vessel away from the vessel wall.

[0051] In contrast, the above described outflow conduit includes a flared portion in which the diameter of the outflow conduit progressively increases towards the outlet, so that the conduit has a larger cross sectional area where the conduit meets the arterial vessel. This widening helps accommodate deformation of the outlet during attachment of the conduit to the arterial vessel, meaning that an effective cross sectional area of the conduit can remain relatively consistent from the heart pump to the arterial vessel, even in light of issues such as stenosis, tissue ingrowth, anastomotic narrowing, or the like. This in turn helps reduce pump afterload and pulsatile inertance allowing the heart pump to operate more efficiently, and reduces jetting, thereby reducing sensitivity to insertion angle, making the conduit easier to implant and reducing adverse effects such as vessel inflammation.

[0052] Accordingly, the above described arrangement therefore minimises flow restrictions and decreases afterload. This allows the heart pump to operate more efficiently, reduces pump power consumption, less battery depletion and exchange rates, and thereby provides a durable and reliable ventricular support device. Decreasing afterload can also help enhance a resting blood flow and / or exercise blood flow of the subject and reduce insertion angle sensitivity.

[0053] As patient outcomes improve, it is likely that there will be more patients, being supported for longer, meaning the ability to achieve these improvements is important in being able to adequately support patients as needed.

[0054] A number of further features will now be described.

[0055] As mentioned above, in addition to reducing the impact of narrowing of the conduit, the outlet 222 is configured to disperse blood flow into the arterial vessel, specifically with the flaring of the outlet operating to reduce the pressure of blood entering the arterial vessel. This restricts blood flow being jetted onto vessel wall and hence assists in preventing damages of the arterial vessel, and allows blood-wall interactions to be reduced, so that losses due to varying viscosity may be minimised.

[0056] In one example, the flared portion increases in internal diameter towards the outlet and an example of this will now be described with reference to Figure 3.

[0057] In this example, the flared portion 222a includes a starting diameter DI and an ending diameter D2. The ending diameter D2 is larger than the starting diameter DI. In one example, the ending diameter D2 is about 1.2 to 1.6 times the starting diameter DI, and is typically about 12 mm to 16 mm, or more typically about 14 mm, depending on the preferred implementation. The starting diameter DI is approximately equal to an internal diameter of the sidewall, and is typically about 8 mm to 12 mm, or about 10 mm.

[0058] The flared portion 222a also has a length L, which in one example is configured to accommodate implantation of the conduit. In this regard, during a heart pump implantation surgery, after positioning the LVAD at the cardiac apex and inserting an inflow conduit into the left ventricle of the heart, the surgeon positions the outflow conduit directing this around the patient’s anatomy until the outlet is proximate the aorta. The surgeon will then align theoutlet so it points into the aorta at an insertion angle, and cut the conduit so that the end of the cut conduit lies substantially parallel with the outer surface of the aorta.

[0059] An example of this is shown in Figure 4, in which the outlet 222 is connected to an arch of the arterial vessel 252. It will therefore be appreciated that the outlet is cut at an angle corresponding to the insertion angle, and so to maintain the flaring of the outlet, the flared portion extends along the conduit so that the outlet is still flared irrespective of an angle of insertion.

[0060] In general, the angle of insertion aims to minimise disruption of laminar flow in the arterial vessel, whilst also accommodating the patient anatomy, and allowing the surgeon access to the perimeter of the outlet to allow this to be joined to the arterial vessel.

[0061] In one example, the outlet 222 may be attached to the arterial vessel at an angle that is less than 90°, between 10° and 70°, about 15°, about 30°, about 45° or about 60° as described in Aliseda A, Chivukula VK, McGah P, et al. LVAD Outflow Graft Angle and Thrombosis Risk. ASAIO J. 2017; 63(1): 14-23.

[0062] To accommodate these preferred angles, the flaring typically extends about 10% to 30% of an outflow conduit length, and typically about 25 mm to 35 mm, or more typically about 30 mm.

[0063] The conduit may also include additional structures to help attach the conduit to the arterial vessel and / or maintain the flared shape. For example, some, or all, of the outlet may be thicker than the sidewall and / or stiffer than the sidewall and / or include at least one surface structure, such as ribbing or the like, to maintain a shape of the outlet. This allows the outflow conduit to be easily attached to the arterial vessel and can help maintain a flared profile upon attachment, for example to avoid kinking or similar.

[0064] In one example, the outlet includes a rim, which can also have a flared profile so that at least part of the flared portion is defined by a flared profile of the rim. This assists in maintaining a shape of the outlet during and / or after attachment.

[0065] In one example, at least part of the sidewall of the outflow conduit is coated. The coating material may be a low friction coating or an anti-biofouling coating, which can assistin further reducing pressure losses in the conduit. Specifically, the low friction coating reduces blood interaction with the conduit sidewall and further reduces the losses due to varying blood viscosity, which minimise pressure loss and thereby improves pump efficiency. Additionally, an anti-biofouling coating can prevent formation of thrombus within the conduit, and can also reduce tissue in-growth into the outlet. In one example, the coating material is a nanocoating material, such as the material described in US20140342954, which creates a liquid repellent surface comprising an anchoring layer, which forms an immobilized molecular anchoring layer on an inner surface of the conduit side wall. The anchoring layer, has a head group that is covalently linked to, or adsorbed onto, the surface and a functional group. The functional group of the treated surface has an affinity for a lubricating layer, which is applied to the treated surface. The anchoring layer and replenishable lubricating layer are held together by non- covalent attractive forces. Together, these layers form an ultra-repellent slippery surface that repels certain immiscible liquids and prevents adsorption, coagulation, and surface fouling by components contained within.

[0066] The conduit can also be made from an at least partially compliant material. In this regard, compliance can help absorb and therefore disperse pressure waves, which in turn reduces the pulsatile inertance. This can in turn reduce excessive pulsatile afterload, thereby decreasing the average flow support provided by the pump and limiting the increase in exercise associated pump flows.

[0067] In one example, the outflow conduit can be made of nanofibers, typically generated from polymers, and in particular natural polymers such as include collagen, cellulose, silk fibroin, keratin, gelatin and polysaccharides such as chitosan and alginate, or synthetic polymers such as poly(lactic acid) (PLA), polycaprolactone (PCL), polyurethane (PU), poly(lactic-co-glycolic acid) (PLGA), poly(3-hydroxybutyrate-co-3 -hydroxy valerate) (PHBV), and poly(ethylene-co-vinylacetate) (PEVA). The nanofibers can be electrospun, and in one example the material used in an electrospun silk.

[0068] However, it will be appreciated that this is not intended to be limiting, and alternative materials such as polyester, polytetrafluoroethylene, polyurethane, biosynthetic materials, or the like could be used.

[0069] In another example, the conduit includes a drug eluting conduit, and in particular is configured to elute a drug that inhibits inflammatory and / or immune system responses that lead to tissue ingrowth. This can further help reduce ingrowth and ensure conduit diameter is maintained.

[0070] A study was performed to examine the frequency, extent and nature of tissue ingrowth within the continuous-flow left ventricular assist device outflow conduits. The study sought to characterise conduit histopathology at explant in a cohort of patients with HeartWare HVAD and assess the effect on pump performance.

[0071] Patients undergoing routine histopathological assessment of a HeartWare HVAD removed at transplant or autopsy were assessed. Outflow conduits were examined macroscopically, and visible tissue was sectioned for microscopic evaluation. In patients who had undergone prior contrast enhanced computerised tomography (CT) with HVAD in situ, the outflow conduit was measured at the aortic anastomosis and 5cm proximal to the anastomosis, in the axial and sagittal planes. All patients had their pump flow, flow pulsatility, current and speed determined from log files examined at 1, 3, 6, 9 and 12 months post LVAD implantation.

[0072] Specifically, stable patients with HVAD (HeartWare, Medtronic) in situ who were undergoing device removal at the time of transplantation or autopsy were included in the study. Ethics approval was obtained from the Hospital’s Ethics Review Committee (Reference 2019 / STE00072). Outflow grafts were examined macroscopically for the presence, extent and location of tissue ingrowth. If present, this tissue was removed from the outflow graft and sectioned for microscopic analysis. Inflammatory reactions were defined as ‘acute’ by the presence of neutrophils; ‘giant cell’ by the presence of multinucleated giant cells and / or granulomas; and ‘chronic’ by the presence of other chronic inflammatory cells. Architectural tissue changes were classified as ‘fibrosis’, ‘neointima’ or ‘atherosclerosis’ at the discretion of the examining anatomical pathologist.

[0073] All patients had their pump flow, flow pulsatility, current and speed retrospectively assessed at 1, 3, 6, 9 and 12 months post-implant from pump controller log files. Pump flow and flow pulsatility were corrected offline for changes in hematocrit. Patients who had undergone contrast-enhanced computerized tomography (CT) had their outflow graft diametermeasured at the aortic anastomosis and 5cm proximal to the anastomosis in the axial and sagittal planes.

[0074] Pump flow changes over time and CT diameter were statistically assessed using the paired t-test or Wilcoxon signed-rank test for normally distributed and non-normally distributed data respectively. A p-value of <0.05 was considered significant.

[0075] 25 consecutive patients were assessed (24 LVAD, 1 BiVAD). One device was assessed at autopsy following a fatal intracerebral hemorrhage. The remainder were assessed following heart transplantation. There were no device explants performed for myocardial recovery or pump malfunction. Twenty one out of 26 (81%) had undergone device implantation via median sternotomy, while the remainder (5 patients, 19%) had undergone a minimally invasive approach via left lateral thoracotomy.

[0076] Of the 26 outflow grafts assessed, there was evidence of tissue ingrowth in 24 (92%). The most common site was the distal anastomosis (18 / 24, 75%), with the graft body involved in 14 / 24 (58%). Microscopic evaluation revealed acute inflammatory infiltrate in 4 / 24 grafts (17%), granulomatous infiltrate in 11 / 24 (46%), and other chronic inflammatory infiltrate in 14 / 24 (58%). The majority of grafts had evidence of either neointima formation (18 / 24, 75%) and / or fibrosis (18 / 24, 75%), with atherosclerosis present in 2 / 24 (8%). The median depth of tissue was 1mm (range 0-2mm).

[0077] Typical microscopic findings show prominent / chronic inflammation, including moderate numbers of plasma cells and occasional neutrophils or some lymphoid aggregates. Additionally, fibrotic tissues were observed with focal mild atheromatous change in addition to a multinucleated giant cell reaction, as well as fibrotic tissue consistent with large vessel intima and scattered chronic inflammatory cells.

[0078] The median time on-pump for those with acute inflammatory infiltrate was 158 days (range 92-327 days), granulomatous infiltrate 238 days (range 92-482 days) other chronic inflammatory infiltrate 306 days (range 92-482 days), fibrotic changes 294 days (range 92- 1032 days) and neo-intima formation 307 days (range 119-1032 days). These differences were not statistically significant. Twelve out of 25 patients (48%) had microbiological evidence of bloodstream (3 / 25, 12%) or driveline (9 / 25, 36%) infection at some stage whilst supported on-pump. There was no association between presence of device-related infection and either inflammatory (p=0.70) or neointimal / fibrotic changes (p=0.61) within the conduit.

[0079] Contrast enhanced CT was performed in 19 patients at a median of 87 days post implant (range 10-296 days) and example images are shown in Figures 5A to 5D.

[0080] In particular, the images show a conduit 523, extending to the aorta 552, with stenosis clearly visible at 560 in Figures 5B and 5D. In particular, these images show a significantly smaller graft diameter at the anastomosis than within the body of the graft, with a mean conduit diameter of 9.5±0.6mm at the aortic anastomosis compared to l l.l±0.5mm 5cm proximal to the anastomosis (p<0.0001).

[0081] There are a number of possible explanations for this stenosis. In particular, the aortic anastomosis was the most common site of tissue ingrowth in patients, and so the CT findings may reflect relative stenosis at this point. Additionally, suturing of the graft is likely to prevent expansion at the anastomotic site, both on a beat-to-beat basis during systole, and due to stretching of graft material in the long-term. Regardless of whether the underlying cause is pathological or a result of outflow graft design and implant technique, the absolute and relative reductions in graft diameter at the anastomosis may have significant effects on both flow rate and turbulence.

[0082] Analysis of patient log files revealed a trend towards decreased pump flow over time (4.61+0.76 vs 4.43+0.80L / min, p=0.07), despite a trend towards increased pump speed (2488+109 vs 2532+125 RPM, p=0.07). Where pump speed remained unchanged compared to baseline, there was a significant (5.8+8.6%) decrease in both pump flow (4.65+0.86 vs 4.38+0.92L / min, p=0.01) and flow pulsatility (5.00+1.10 vs 4.16+1.05L / min, p=0.006) over a median time period of 5 months (range 2-11 months).

[0083] These results demonstrate there is evidence of tissue formation within the HVAD outflow conduit in the vast majority of patients, most commonly located at the aortic anastomosis. This is associated with significantly decreased pump flow over time.

[0084] With approval of cf-LVAD devices for destination therapy in severe heart failure, a greater understanding the long-term pump-patient interaction, including the effects of the pumpon the body, but also the effects of the body on the pump, is required. It is well recognized in the non-LVAD setting that synthetic vascular graft insertion can cause a foreign body reaction and neointima formation, with a demonstrated association between these processes and graft failure in one analysis of arteriovenous fistulae. In the setting of cf-LVADs, meanwhile, case reports have identified stenosis at the aortic anastomosis as a rare but important cause of pump dysfunction. The current study systematically assesses the human tissue response to the presence of the cf-LVAD outflow graft, and long-term changes in pump flow, in a consecutive cohort of patients without evidence of pump dysfunction or clinical outflow obstruction.

[0085] The key findings of the study are: firstly that some degree of tissue ingrowth into the outflow graft is a near-universal phenomenon, and most commonly occurs at the aortic anastomosis; secondly, that this reaction may variably be comprised of an acute inflammatory, chronic inflammatory, fibrotic or neointimal reaction; and thirdly, that this is associated with a measurable decrease in pump flow and flow pulsatility over time.

[0086] It is therefore apparent that outflow graft characteristics have a significant impact on the function of rotary pumps through afterload effects. It has been demonstrated that hemodynamically significant gradients are generated across the HVAD outflow graft in a pulsatile circulation; furthermore, minimum luminal area has a significant, non-linear effect on the magnitude of these gradients. From the baseline HVAD conduit cross-sectional area of 0.79cm2- smaller than the widely accepted threshold for severe aortic stenosis - a circumferential tissue ingrowth depth of 1mm would result in a 36% reduction to 0.50cm2. This would be expected to be sufficient to cause a significant reduction in pump performance due to increased pressure head across the cf-LVAD impeller. This is reflected in the results of the study, in which there were statistically significant reductions in pump flow and flow pulsatility over time where pump speed was kept constant.

[0087] Although the absolute magnitude of pump flow reduction across the whole cohort was small, the upper range limit of 0.95L / min is a clinically significant decrease. Furthermore, the observed trend towards increased pump speed over time lends further weight to the hypothesis that, in some patients at least, the reduction in pump flow over time may be significant enough to warrant clinician intervention.

[0088] The precise characteristics of the tissue response were highly variable in terms of extent, location and histopathological features; moreover, the factors that influence these characteristics are unclear. There was no association between time on-pump and depth of neointima assessed histopathologically. This finding is not unexpected, given the results of a previous large animal study assessing neointima formation within synthetic coronary grafts, which showed that following an initial period of endothelial and smooth muscle cell hyperplasia lasting 3 months, a ‘steady -state’ is achieved in which there is turnover but not further accumulation of neointima.

[0089] Additionally, neither time on-pump nor the presence of driveline or bloodstream infection significantly affected the location of tissue ingrowth, or whether or not this tissue was inflammatory in nature.

[0090] In any event, it is proposed that these long-term effects of tissue ingrowth on pump performance could be mitigated using the above described conduit design. Specifically, the conduit diameter is increased at least a distal end, so that the outlet of the conduit is flared. Given the non-linear effect of graft diameter on pressure losses, the increased diameter means that the same degree of diameter reduction would result in a smaller increase in pressure gradient and therefore a smaller reduction in pump performance.

[0091] To further demonstrate the effect of conduit diameter on pump outflow, measurements were performed of pressure gradient losses along the length of a tubular outflow conduit to a point of aortic insertion, for conduits having 8, 10 and 12mm diameters, at pump speeds of 3000 and 2600 rpm. Results are shown in Figure 6, which highlight that there is little difference between 10 and 12mm, with a major difference for an 8mm conduit. This reinforces the view that larger diameter conduit and flared outlet (to reduce the impact of tissue growth), will result in improved pump outflow.

[0092] Further, a computational fluid dynamics study was performed to examine outflow into the aorta for different outflow conduits, with Figures 7A to 7C showing resulting flow lines for 8, 10 and 12mm diameter conduits. In this example, the 12mm conduit is marginally flared. The modelling shows improved diffusion in the aortic root and into the major cerebral vessels, with reduced outflow resistance leading to improved outflow.

[0093] A further potential modification would involve the use of drug-eluting grafts. For example, both Paclitaxel and Sirolimus-eluting grafts have demonstrated efficacy in preventing synthetic arteriovenous fistula restenosis in animal studies, although they are not yet marketed for human use.

[0094] Throughout this specification and claims which follow, unless the context requires otherwise, the word “comprise”, and variations such as “comprises” or “comprising”, will be understood to imply the inclusion of a stated integer or group of integers or steps but not the exclusion of any other integer or group of integers. As used herein and unless otherwise stated, the term "approximately" means ±20%.

[0095] It must be noted that, as used in the specification and the appended claims, the singular forms “a,” “an,” and “the” include plural referents unless the context clearly dictates otherwise. Thus, for example, reference to “a support” includes a plurality of supports. In this specification and in the claims that follow, reference will be made to a number of terms that shall be defined to have the following meanings unless a contrary intention is apparent.

[0096] It will of course be realised that whilst the above has been given by way of an illustrative example of this invention, all such and other modifications and variations hereto, as would be apparent to persons skilled in the art, are deemed to fall within the broad scope and ambit of this invention as is herein set forth.

Claims

THE CLAIMS DEFINING THE INVENTION ARE AS FOLLOWS:1) An outflow conduit for a heart pump, the outflow conduit including: a) an inlet configured to be attached to a pump outlet of the heart pump to receive blood therefrom; b) an outlet configured to be connected to an arterial vessel, wherein the outlet includes a flared portion to minimise flow restrictions; and, c) a sidewall defining an internal cavity extending from the inlet to the outlet, the cavity being configured to transport blood from the heart pump to the arterial vessel.2) The outflow conduit according to claim 1 , wherein the outlet is configured to disperse blood flow into the arterial vessel.3) The outflow conduit according to claim 1 or 2, wherein the flared portion increases in internal diameter towards the outlet.4) The outflow conduit according to any one of claims 1 to 3, wherein the flared portion includes a starting diameter and an ending diameter, wherein the ending diameter is larger than the starting diameter.5) The outflow conduit according to claim 4, wherein the ending diameter is at least one of: a) about 1.2 to 1.6 times the starting diameter; b) about 12 mm to 16 mm; and, c) about 14 mm.6) The outflow conduit according to claim 4 or 5, wherein the starting diameter is at least one of: a) approximately equal to an internal diameter of the side wall; b) about 8 mm to 12 mm; and, c) about 10 mm.7) The outflow conduit according to any one of claims 1 to 6, wherein the flared portion is at least one of: a) about 10% to 30% of an outflow conduit length; b) about 25 mm to 35 mm in length; and, c) about 30 mm in length.8) The outflow conduit according to any one of claims 1 to 7, wherein the outlet is configured to be connected to an arch of the arterial vessel.9) The outflow conduit according to any one of claims 1 to 8, wherein the outlet is at least one of: a) thicker than the sidewall; and, b) stiffer than the sidewall.10) The outflow conduit according to any one of claims 1 to 9, wherein the outlet includes at least one surface structure to maintain a shape of the outlet.11) The outflow conduit according to any one of claims 1 to 10, wherein the outlet includes a rim.12) The outflow conduit according to claim 11, wherein the flared portion is defined by a flared profile of the rim.13) The outflow conduit according to any one of claims 1 to 12, wherein at least part of the sidewall is coated.14) The outflow conduit according to claim 13, wherein coating is at least one of: a) a low friction coating; and, b) an anti-biofouling coating.15)The outflow conduit according to any one of claims 1 to 14, wherein the conduit is made of an at least partially compliant material.16) The outflow conduit according to claim 15, wherein the conduit is made from at least one of: a) nanofibers; b) electrospun nanofibers; c) silk; d) electrospun silk; and, e) electrospun silk with a nanocoating.17)The outflow conduit according to any one of the claims 1 to 16, wherein the conduit includes a drug eluting conduit.18) The outflow conduit according to claim 17, wherein the drug inhibits at least one of: a) inflammatory response; and, b) immune system responses.

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