Section 1: Introduction
The heart-lung machine (HLM) is the device that has allowed modern open-heart surgery to be conducted with relative ease while maintaining sufficient oxygenated blood in the patient’s body [1]. The machine temporarily replaces the role of the heart and lungs, allowing surgeons to conduct their surgeries on a ‘non-beating’ heart and provides them with a bloodless field. This replacement process is known as a cardiopulmonary bypass (CPB) [2]. The HLM is used in several different cardiac surgeries, including valve repairs and replacements, repairing heart defects and revascularizing blocked arteries. By performing these surgeries on non-beating hearts, cardiothoracic surgeons are able to ensure greater precision and fewer complications.
Originally developed as a perfusion pump by Dr Alexis Carrel and Charles Lindbergh in the year 1930, the medical device has been modified over 20 years [3]. Later, in 1953, Dr John H. Gibbons performed the first successful surgery (repairing an atrial septal defect) using a HLM [4]. Today, the HLM is almost an essential component of a cardiothoracic operation theatre (OT). The first open-heart surgery supported by a HLM in Singapore was conducted in 1965 at the Singapore General Hospital [5]. Since then, the use of the HLM in hospitals in Singapore has grown. Over the years, several medical device companies have tried to improve the designs of the parts of the HLM, trying to make it as efficient and error-free as possible.
This report will cover the basic functioning of the HLM, the functions of the different components, the issues generally found in the machine, solutions that have been developed over time and comparison of a few different brands.
Section 2: About the Heart-Lung Machine
The HLM has a very specific working manner and each part has its own importance. This section will discuss how the heart-lung machine works and what the function of the major components are. Companies that create HLMs choose to improve different components to make their products the best in the market. This will also be discussed in this section.
2.1 How does the heart-lung machine work
During a CPB, deoxygenated blood is transferred from the veins of the patient into the blood reservoir in the machine through a silicone rubber tubing. This is where the blood will temporarily be stored. Blood is then transported via a pump, through a heat exchanger. The heat exchanger monitors, controls and maintains the blood temperature at the optimum level; less than 37°C. The blood then moves into an oxygenator that contains a partially permeable membrane. This membrane allows carbon dioxide to diffuse and move out of the bloodstream and allows oxygen to diffuse into the bloodstream. This oxygenated blood is then transported back into the patient’s body, into the arteries, supplying cells with the required oxygen even during the open heart procedures.
Fig 1: Visualisation of the working mechanism of the heart-lung machine
2.2 Components of the machine and their functions
The HLM has several components; pumps, cannulae, tubing, a reservoir, oxygenator and an arterial line filter. Apart from these, certain machines also have systems to monitor pressures, temperature, oxygen saturation. haemoglobin, blood gases and electrolytes. Some machines also come equipped with safety features like bubble detectors, oxygen and reservoir low-level detection alarms [2].
The arterial pump is an important component of the HLM and can be either a roller pump or a centrifugal pump.
The roller pump acts as a positive displacement device that allows for a constant delivery of blood volume. The PVC or silicone tubing carries blood between two rollers and gets compressed between the rollers. These compressions create variations in blood pressure at the outlet. The blood flow in the pump is dependent on the internal diameter of the pump tubing, rotation rate of the rollers and diameter of the pump head.
The centrifugal pump has vaned impellers magnetically coupled at its base using an electric motor. These create a centrifugal force when rotated at high speeds. The centrifugal force is converted to kinetic energy and potential energy, causing blood flow and pressure, respectively [2].
Cannulae connect the patient to the HLM via a circuit. They are made of polyvinyl chloride (PVC) and are reinforced by wires to prevent kinking in the tubing that can cause obstructions. PVC is used in the cannulae and other tubings for the machine because of its durability and acceptable haemolytic rate (rate of destruction of blood cells) [2].
Oxygenators ensure and regulate oxygen content in the blood that flows through the machine. Previously, bubble oxygenators were more commonly used. HLM nowadays primarily use membrane oxygenators. These consist of microporous polypropylene fibres (100-200 μm internal diameter). Blood flows outside the fibre, separating the blood and gas phases. Membrane oxygenators reduce the occurrence of air emboli and give greater accuracy in controlling the blood gases. The oxygenator is usually coupled with a heat exchanger which is placed proximally to the oxygenator reduce the release of gaseous emboli (blockage causing material) due to changes in the temperature of oxygen-saturated blood [2].
2.3 Comparison of alternatives in the market available in Singapore
Two of the most prominent medical device brands in the market producing HLMs are Medtronic and LivaNova. Both the companies are successfully making more than one billion USD from the devices that they have created. There are mainly three components that need to be compared. First, the centrifugal blood pump. This is one of the most important components in the HLM. It is required to transfer blood from the body to the device and back to the body. A good blood pump aims to minimise blood clots due to erythrocyte damage. The Medtronic ‘Affinity CP’ and LivaNova ‘Revolution’, have higher flow rates of nine litres/min [7] and 8 litres/min [8] respectively. The Medtronic has a low priming volume of 40ml [7] while LivaNova has approximately 57ml of priming volume [8]. When the blood flow is too high, red blood cells (RBC) will get damaged due to high shear [9]. Also, when more blood is needed to prime the device, the greater the chances for a blood clot to occur. This is because blood may clot in the device and the tubes which connects the device and the patient. Overall, the shape and size of the blood pumps used in LivaNova and Medtronic are about the same. However, another company, Abbott Laboratories, has created a better version of the blood pump in terms of technology. It utilises the Full Maglev flow technology to minimise blood clots in the device. The rotor is levitated in the blood pump and the rotational speed is precisely controlled [10].
The oxygenator system is the second most important part of the HLM as it is required for gas exchange between oxygen and carbon dioxide. The oxygenator system in the LivaNova ‘Inspire’ has a flow rate of six litres/min [11] while the Medtronic ‘Affinity Fusion’ has a flow rate between one to seven litres/min [12]. A high flow rate will result in the damage of the RBC. In addition, the difference between the two versions is the presence of Trillium Biosurface or Cortiva BioActive Surface in the Medtronic version. These surfaces helps to prevent conditions such as systemic inflammatory response syndrome that affects the heart, lungs, brain and other organs [13].
Overall, the S5 Model has better integration with the perfusion circuit and allows for a quick and easy exchange of individual components, making sure the machine works properly. Since upgrades can be implemented at any time, the S5 system will cater to all future devices as well. The system also has a lot of safety features in place. All the warnings and automated alarms of the system are recorded and listed chronologically on the system display panel. Information and incidents cannot get lost, and critical situations can be quickly detected and cleared. There are text messages for alarms and warnings with differentiated alarm tones. It also has bubble detectors, level controls and pressure controls. Previously saved data can be moved to the new S5 system using an adapter cable. Additionally, the flexibility of the system allows the system to be moved easily to optimise its layout in smaller OT’ as well and suit the preferences of the perfusionist. The perfusionist is also able to access all the controls at ease. System panel can be easily mounted and adjusted. The care taken in the selection of the materials and parts used helps to minimize the operating costs and long-term costs of the system.
The Medtronic Performer CPB model caters to the present spatial challenges in operating rooms with less OT space utilised whilst allowing convenient access for perfusionists. With disposables at operating-table height and monitoring, pumping, controlling, filtration, air handling, or cardioplegia functions available when required, the Performer CPB, reduces the morbidity of cardiac surgery markedly. System commands support specific pump function, and a central information display with colour touch-screen integrates graphic menus and buttons for user-selectable preferences. While connected to the machine, the system is able to monitor all physiological functions of the patient. With its audio and visual features like the Air Bubble Detection, Active Air Removal technology, oxygen saturation and haematocrit monitoring and level sensors to monitor changes in the fluid in reservoirs, any unusual physiological changes can immediately be tended to. With an alterable height, the instrument can also be modified to suit the patient. The system also contains data management functions to capture, display, store and transmit all key instrument-related perfusion parameters. Data in the system can be transferred to an independent Microsoft(R) Office Excel-based macro program for report and graph generation and case records, enabling coordination of all of the monitoring devices [22][23][31].
Section 3: The key medical problems and inefficiencies associated with the Heart-Lung Machine
Due to its mechanical components and their interaction with blood, HLM can produce significant effects on the body. Factors such as blood damage, difficulty in determining appropriate proportion of gases transferred in oxygenator, and contact between blood and artificial materials can give rise to problems that arise during or after surgery in the intensive care unit [17].
3.1 Inefficiency of Blood Pump
Hemolysis
Hemolysis arises from three situations: naturally by the spleen, physico-chemical imbalance, or through exposure to non-physiological mechanical stress [14]. In HLMs, hemolysis occurs mechanically, primarily in the blood pumps, either from the direct trauma of blood passing through the rollers or upon exposure to different surfaces at varying speeds [14]. The latter refers to shear stress created in the blood pump experienced by the RBC, which destroys them. Hemolysis is present in blood pumps, as shown in various studies by the increasing levels of plasma-free haemoglobin (PfHb) and decreasing levels of haptoglobin during and after CPB [15].
Evidence from studies have indicated that blood circulated in HLM undergoes hemolysis. These studies use markers that detect hemolysis [15]. An example is haptoglobin concentration. It is a protein that forms bonds with haemoglobin and produces a complex that restricts renal loss of haemoglobin, thus reducing its quantities in the bloodstream. Therefore, a decrease in haptoglobin concentration indicates hemolysis [17].
Hemolysis is likewise indicated by elevated levels of plasma-free haemoglobin (PfHb) in the plasma. As the RBC bursts, PfHb is released into the bloodstream before being eradicated by Hb scavengers, such as haptoglobin. However, due to decreased haptoglobin levels, there is reduced eradication of PfHb, causing higher PfHb levels. Data from various studies coherently display elevated plasma Hb concentrations and reduced haptoglobin levels during CPB, indicating RBC destruction [18]. One study even displayed delayed hemolysis after CPB [18].
Shear Stress
Blood undergoes hemolysis because of shear stress. It is defined as a force applied to an area of liquid confined between two plates. The velocity at which the liquid moves in proportion to the separation of the plates is known as the velocity gradient, or shear rate. This velocity gradient comes about due to layers of fluid moving at different velocities between 2 plates as can be seen in Fig 1. Shear stress is related to shear rate with the following equation (Eq. 1).
Shear stress = Viscosity x Shear rate [Eq. 1]
Fig 1. The bottom plate is stationary, whilst the top plate is moving, hence, creating a velocity gradient due to the layers of fluid moving at different velocities between the 2 plates
With this understanding, shear stress exists in blood pumps since the housing wall is stationary, whilst the impeller wall is moving. There is greater shear stress at the stationary wall as compared to the moving impeller wall. Hence, RBC flowing through blood pumps is caught in a shear flow with one end of the cell dragged by the rotating impeller to move at high velocities while the other end is dragged to stay put by the stationary housing wall, causing deformational stresses on the RBC that may cause it to burst.
This is further supported by studies done on wall shear stress in centrifugal and roller pumps. Fig 2. shows the shear stress distribution in centrifugal pumps. As can be seen, there are regions with shear stress greater than 150Pa which is the threshold shear stress of RBC, which can cause hemolysis of RBC.
Fig 2. The distribution of shear stress in centrifugal blood pump.
Studies made on blood damage using a rotational viscometer show that there is a threshold shear stress, 150Pa, above which extensive cell damage is directly due to shear stress. At high stresses, very high rates of hemolysis occur [20]. RBCs have a high tolerance to most forces (e.g. wall impact forces, positive and negative pressure)[15], except for shear stress. Acceptable values of shear stress range from 0.01Pa to 0.5Pa. From 150Pa onwards, damage to RBCs occur [15].
These studies show that shear stress results in RBC damage when the shear stress threshold of RBC (150Pa) is exceeded. As there exists high shear stress in regions of blood pumps as discussed earlier, this proves that blood pumps cause hemolysis. Thus, layers of fluid moving in different velocities in medical cardiovascular devices result in high shear stresses that cause hemolysis.
3.2 Inefficiency of Membrane Oxygenator
Three variations of blood oxygenators have been manufactured for use in HLM. Presently, oxygenators administer gas exchanges to blood through a membrane (sheet or hollow-fibre oxygenators). Sheet membrane oxygenators are structured as flat plate microporous polypropylene membranes or spiral wound silicone membranes.
Firstly, microporous membranes comprise of pores of length 0.05 to 0.3 µm, which produces a direct blood gas interface till a fine protein layer rapidly materialises, creating molecular membranes (Fig. 9). Silicone membranes, upon lengthy durations of CPB, resemble microporous membranes, and are recognised as ‘true membranes’ and ‘plasma leakage’ inhibitors. Plasma leakage results from a loss of hydrophobicity through adsorption of protein by the micropores. Plasma is discharged from the micropores; thus, gas diffusion decreases [24]. Hollow-fibre oxygenators possess larger surface area-to-blood volume ratios and comprise predominantly of microporous polypropylene or polymethylpentene material.
In the human lungs, the physical processes of convection, diffusion and chemical reactions that take place during gas exchange mirrors that in the artificial lungs but less competently. Diffusional distances are greater, approximately 200mm (10 mm in the human alveolus), and surface area for gas exchange is 1.7–3.5 m2 compared with 70–100 m2 in the human lung.
Gas exchange obeys Fick’s law of diffusion:
where: Vgas is the quantity of gas transferred
A is area
D is a diffusion constant
ẟP is the partial pressure difference
T membrane thickness
Sol specific gas solubility
MW molecular weight.
By application of Fick’s law to the mechanical lung, a larger difference in partial pressure of oxygen (maximum pressure gradient of 760 mm Hg minus oxygen tension in blood) and to a certain degree, carbon dioxide (maximum gradient equals to carbon dioxide tension in the blood phase) is needed to offset greater diffusional distance and controlled, minimised surface area. Although the pressure gradient for carbon dioxide (CO2) is significantly reduced compared to oxygen (O2), the gas permeability of mechanical lung element is higher to CO2 than O2; silicone membranes possess a CO2 : O2 diffusion ratio of approximately 5 : 1 [26].
Oxygen transfer in modern blood oxygenators is described by the advancing front theory. Based on research on oxygen movement through constantly shifting blood films, two specific sections in the blood film are formed due to the substantial binding ability of haemoglobin for oxygen and rapid oxygen– haemoglobin kinetics. Oxygen diffuses via a completely oxygen saturated section to an unsaturated front where the oxygen bonds with unsaturated haemoglobin [4].
Hence, speculation of carbon dioxide transfer in modern oxygenators is complicated, due to factors like dissolved CO2, plasma bicarbonate ions and varying partial pressure of CO2 in the gas phase.
3.3 Systemic Inflammatory Response
During CPB, blood travels through the polyvinyl chloride tubing, which is used for the blood to flow in HLM. Since the blood is not flowing through the blood vessel which it used to flow in, the blood will naturally react abnormally, as it is trying to protect the body when it comes in contact with foreign objects [28]. Therefore, despite using material which is compatible with the blood to make the tubing, the blood still reacts abnormally to it, leading to system inflammatory response syndrome (SIRS). This usually happens after CPB, when the body reacts to the new object present in the blood. The body will react as such: body temperature above 38°C or below 36°C, heart rate above 90 beats per minute, breath rate above 20 breaths per minute, or leukocyte above 12000 [29].
Another factor such as ischemia-reperfusion is also a cause of SIRS. This happen when the vital organs such as the lungs, brain, heart, and kidney did not receive enough oxygen during the CPB process. The inflammatory response, which leads to SIRS, comes after the ischemia period [30].
Section 4: Improvements to Heart-Lung Machines
With the advancements in technology since 1845, the components of the HLMs have continuously been improved to tackle their inefficiencies, giving rise to the modern-day HLMs. The following portion will address some of the improvements made along the years, such as the spiral model of centrifugal pumps, silicone membranes, MicroMox, and the heparin-coated internal CPB circuit.
4.1 Blood Pumps
Both roller pumps and centrifugal pumps cause damage to blood due to shear stress. The following are possible improvements in HLM to combat the inefficiencies of its primary component, the blood pump.
Firstly, the roller pump is still the most commonly used blood pump. To improve hemolysis rates in HLM using roller pumps, some hospitals switch to centrifugal pumps to eliminate the intermittent tubing occlusion. Centrifugal pumps have lower hemolysis rates as studies have shown [32][33][34] as roller pumps produce great negative pressure when pulling blood into the CPB circuit, which can result in high hemolysis rates [34]. They are also safer as they have an added safety feature. The flow generated by the centrifugal pump inversely depends on the pressure in CPB circuit. Hence, when the peripheral resistance of the patient rises, blood pressure also rises, causing pressure in CPB circuit to rise, resulting in reduced flow. This prevents blood pressure in patients from building up too high, which may cause further complications [35].
For HLMs using centrifugal pumps, improvements have been made to improve the hemolytic rate such as the development of a new model that uses a spiral impeller. This new model concurrently applies centrifugal and axial pumping principles, thus enhancing pumping performance with no rise in hemolysis rates. Axial pumping effect is generated due to double entrance threads at the impeller surface (Fig 3) [36]. Centrifugal pumping effect is generated because of the impellers’ conical shape (Fig 3) [36]. Combination of axial and centrifugal pumping principles generates improved hydrodynamic performance without elevating hemolysis rates as shown in studies, where hemolysis rates in this new spiral model are assessed. These studies, as opposed to research on previous models, have shown improved results [36][37][38].
Fig 3. Spiral pump model
4.2 Membrane Oxygenator
In the case of ‘plasma leakage’, the use of silicone membranes as opposed to microporous membranes in the oxygenator tackles the problem (Fig. 11). Differing partial pressures of CO2 in the gas phase can be reduced through increasing the ventilating gas flow rate (sweep gas) entering the oxygenator gas phase [40].
Lastly, in 2011, research was conducted in which a modified Computational Fluid Dynamics Model was established and substantiated through use of a uniquely created micro membrane oxygenator (MicroMox). The MicroMox was constructed such that the fibre distribution and cluster geometry were highly reproducible and possible flow channeling is prevented. Its minuscule size (VFluid = 0.04 mL) makes reproduction of the cluster of 120 fibres possible. The simulation liquid used was a non-Newtonian blood substitute. Physical solubility and chemical bonding of O2 and CO2 in blood was defined through a theoretical model. Fixed oxygen partial pressure at the fibre pores and a constant state flow field were factors in the calculation of the mass transmission. To resolve the MicroMox fibre bundle, in the direction of flow, mass transport was prompted for symmetrical geometric sections. This was validated by measuring the in vitro gas transfer rates of the MicroMox. These measurements were performed in accordance with DIN EN 12022 (2) using porcine blood. The theoretical results for mass transfer for two different fibre arrangements coincided with experimental data obtained for different mass flows and constant inlet pressures. Hence, a validated model for the theorised gas exchange in the hollow fibre oxygenators could be determined
4.3 Inflammation Response
One way to alleviate inflammation response is to coat the internal CPB circuit using heparin. With a heparin coated circuit, the circuit is more biocompatible for the blood flow [40]. This coated surface also makes the CPB circuit water and protein resistant and promotes lipoprotein binding. In addition, it causes leukocytes and endothelial cells to release lesser toxic mediators, causing an increase in static lung compliance, and a decrease in both, alveolar-arterial gradient and resistance in pulmonary vascular [41]. However, this favourable effect is usually observed at the end of surgery and does not reduce the patient’s stay under intensive care. Since heparin coating enhances thromboresistance, a smaller volume of heparin can be injected into the patient’s body for surgery, reducing bleeding and blood transfusion after operation [35].
4.4 Future Innovation
There are several studies currently investigating the combination of interventional and minimally invasive methods for cardiothoracic surgeries. Percutaneous closure devices and robotic technologies like the Da Vinci Robot are examples of these. These innovations reduce complications, blood usage, infections, postoperative pain and recovery time. Therefore, they can be implemented in place of HLMs in hospitals.
Section 5: Conclusion
HLM might have revolutionised the way cardiovascular surgeries are conducted since the early 1950s. However, with advancements in technology, several inefficiencies related to HLM have been identified such as blood damage caused by blood pumps, inefficiencies in the oxygenator, systemic inflammation and arrythmia, to name a few. Over the years, countless developments have been implemented to improve the HLM and create the models we see today. These improvements such as the use of spiral blood pumps, heparin coated circuits, do boost mortality rates of patients who require the use of HLM during surgery. However, these inefficiencies cannot be completely eradicated. There are still some drawbacks to the use of HLM. For instance, studies have shown that the mortality rates for patients who undergo CPB is 1.6 times higher during the first 30 days after surgery [42]. Therefore, we feel that the best solution is to avoid the use of HLM as much as possible, especially since there are other new innovative methods of cardiovascular surgery being introduced.
This reduces the adverse effects experienced by patients who undergo CPB, increasing the lifespans of patients who undergo cardiac surgery. However, in unavoidable circumstances, we believe that based on our comparisons, Medtronic is a more reliable brand to choose.
In conclusion, we feel that Medtronic’s Performer CPB is a much better model as it has better safety features in place as discussed above. The Performer model also uses an oxygenator with Cortiva BioActive Surface which help prevent systemic inflammatory responses in the patient. Even though the LivaNova S5 may be more cost effective in the long run as it can upgraded and repaired part by part instead of buying a whole new machine, we feel that safety measures trump the costs of the machine. In the line of healthcare, life is priceless and patient care is of utmost importance. Hence, therapies that are of value, prolongs life, and alleviates suffering are vital. No amount of money spent on healthcare is deemed wasted.