Biomedical Engineering - Biomachines

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Complete course

Politecnico di Milano SCUOLA DI INGEGNERIA INDUSTRIALE E DELL’INFORMAZIONE Laurea Magistrale – Ingegneria Biomedica Dispensa di Biomachines Docente Prof. Francesco DE GAETANO Studente Loris BARBIERO – 976048 Anno Accademico 202 1 – 202 2 pag. 2 pag. 3 Sommario Sommario ................................ ................................ ................................ ................................ .......................... 3 FLUID MECHANICS SUMMA RY ................................ ................................ ................................ .......................... 4 Fluid viscosity ................................ ................................ ................................ ................................ ................. 4 Fluid mechanics ................................ ................................ ................................ ................................ ............. 5 Bernoulli’s principle ................................ ................................ ................................ ................................ ....... 6 Hagen –Poiseuille Law ................................ ................................ ................................ ................................ ... 8 She ar stress in a cylindrical conduct ................................ ................................ ................................ .............. 8 THE BLOOD ................................ ................................ ................................ ................................ ........................ 9 Composition and function ................................ ................................ ................................ ............................. 9 Red blood cells ................................ ................................ ................................ ................................ ............. 10 Blood viscosity ................................ ................................ ................................ ................................ ............. 11 Coagulation ................................ ................................ ................................ ................................ .................. 12 Haemo lisys ................................ ................................ ................................ ................................ ................... 15 CARDIAC PHYSIOLOGY ................................ ................................ ................................ ................................ ..... 17 Anatomy of the heart ................................ ................................ ................................ ................................ .. 17 Cardiac output (CO) ................................ ................................ ................................ ................................ ..... 18 Frank -Starling mechanism ................................ ................................ ................................ ........................... 19 Guyton model ................................ ................................ ................................ ................................ .............. 20 HEART VALVE PROSTHESIS ................................ ................................ ................................ .............................. 24 History ................................ ................................ ................................ ................................ ......................... 25 Determination of the optimal diameter of a pr osthetic heart valve ................................ .......................... 30 EXTRACORPOREAL CIRCULATION ................................ ................................ ................................ .................... 34 Cardio Pulmonary Bypass ................................ ................................ ................................ ............................ 35 ECC parameters ................................ ................................ ................................ ................................ ........... 37 Reservoir ................................ ................................ ................................ ................................ ...................... 42 Pumps ................................ ................................ ................................ ................................ .......................... 43 Oxygenator ................................ ................................ ................................ ................................ .................. 52 Heat Exchanger ................................ ................................ ................................ ................................ ............ 59 Filters ................................ ................................ ................................ ................................ ........................... 61 Cardioplegia ................................ ................................ ................................ ................................ ................. 65 pag. 4 FLUID MECHANICS SUMMARY Fluid viscosity By definition, fluid viscosity is responsible for «internal friction»; it is a measure of its resistance to gradual deformation when shear stress is applied. According to the International System of Units (SI), the viscosity is defined as: Accordingto the centimetre –gram –secondsystem of units (cgs), on the other hand, the viscosity is defined as: The shear stress: - Is zero when the fluid stops moving; - Is applied in the opposite direction to the fluid motion; - Has a magnitude directly related to the viscosity of the fluid and the shear rate. Depending on their viscosity behavior as a function of shear rate, fluids are characterized as Newtonian or non -Newtonian. Newtonian fluids have a simple linear relation between shear stress and shear rate. On the other hand, non -Newtonian fluids display a non -linear relation between shear stress and shear rate, have a yield stress, or viscosity that is dependent on time or deformation history (or a combination of all the above). Shear rate dependency on fluid type and applied shear stress: pag. 5 The power’s law fluid relationship is the most common generalized non -Newtonian fluid law for which the shear stress is given by: Fluid mechanics Generally, in the human body, blood flow is LAMINAR. However, under high flow conditions, particularly in the ascending aorta, laminar flow can be disturbed becoming TURBULENT. Turbulence increases the energy required to drive blood flow in the vessel due to the loss of energy in the form of frictions, which generates heat. Reynolds number is a criterion of whether fluid flow is steady (laminar) or on the average steady with small unsteady fluctuations (turbulent). Whenever the Reynolds number is less than about 2’000 , flow in a pipe is generally laminar, whereas, at values higher than 4’000 , flow is usually turbulent. The transition between laminar and turbulent flow does not occur at a specific value of the Reynolds number but in the range 2'000 to 4’000. The Womersley number is a dimensionless number, which is an expression of the pulsatile flow frequency in relation to viscous effects. The Womersley number is essential in order to keeping dynamic similarity when an experiment is scaled up (in particular for experimental study in the vascular system). pag. 6 Bernoulli’s principle The Bernoulli’s principle is a specific case of the energy conservation. Let’s start considering a condu ct with a fluid flowing inside. =f we don’t take into account the thermal losses and the mechanical work, we can write: In particular ������is the sum of kinetic energy, potential energy and flow work (pressure -volume work). Under the hypothesis of steady state: BUT if we consider the “real world” we have to take into account: - Frictional losses( ������≠0); - Local Losses (curvature, valve, variation of section, etc). pag. 7 Accordingto the Darcy -Weisbach equation, the hydraulic slope is defined as: Experimental data about Darcy friction factor have been collected for a wide range of roughness and Re numbers in the MOODY CHART pag. 8 Hagen –Poiseuille Law In 1839 Hagen, and subsequently in 1940 Poiseuille, demonstrate that for a smooth, horizontal, cylindrical and rigid tube with a laminar flow field, the Bernoulli Equation can be simplified as below: Shear stress in a cylindrical conduct The general equation describing wall shear stress in a cylindrical conduct is: for a laminar flow: pag. 9 THE BLOOD The term haemodynamic describes the physical factors governing blood flow within the circulatory system. Blood flow (Q) through an organ is determined by the perfusion pressure (ΔP) driving the flow divided by the resistance (R) to flow Composition and function The blood passing through the cardiovascular system (tissue in a form of a liquid in which there are suspended cells) consists of the following formed elements: - Platelets (1.5÷4)·10 3/mm 3 flat disc -shape (haemostasis function, 5% of RBC); - White blood cells (WBCs) (4÷10)·10 3/mm 3 spherical shape (immunedefence mechanism); - Red blood cells (RBCs) or erythrocytes 5·10 6/mm 3 biconcave disc -shape; - Plasma Blood is a fluid connective tissue that circulates through the arteries to reach the body’s tissues and then returns to the heart through the veins. When blood is “spun down” in a centrifuge tube, the RBCs precipitate to the bottom of the tube, where they account for about 45% of the blood volume. This is called the haematocrit and normally ranges from 40% to 50% in males and 35% to 45% in females. The next layer is a “b uffy coat,” which makes up slightly less than 1% of the blood volume and includes WBCs (leukocytes) and platelets. The remaining 55% of the blood volume is the plasma (serum is plasma with the clotting factors removed and includes water, plasma proteins, a nd various solutes, if this portion is red it means that there has been a massive hemolysis and this can bring to the death of the patient). pag. 10 The functions of blood include: - transport of dissolved gases, nutrients, metabolic waste products, and hormones to and from tissues - prevention of fluid loss via clotting mechanisms - immune defense - regulation of pH and electrolyte balance - thermoregulation through blood vessel constriction and dilation Red blood cells The erythrocyte, commonly known as a re d blood cell (or RBC), is by far the most common formed element: a single drop of blood contains millions of erythrocytes and just thousands of leukocytes. Specifically, males have about 5.4 million erythrocytes per microliter (μL) of blood, and females ha ve approximately 4.8 million per μL. =n fact, erythrocytes are estimated to make up about 25 percent of the total cells in the body. They are quite small cells, with a mean diameter of only about 7 –8 μm . The primary functions of erythrocytes are to pick up inhaled oxygen from the lungs and transport it to the body’s tissues, and to pick up some (about 24 percent) carbon dioxide waste at the tissues and transport it to the lungs for exhalation, to do this they remain within the vas cular network. The RBCs are biconcave disks: they are plump at their periphery and very thin in the center. Since they lack most organelles, there is more interior space for the presence of the hemoglobin molecules that transport gases. Inside the RBCs the re are also electrolytes, oxygen, carbon dioxide, plasmatic liquid. The free hemoglobin is the hemoglobin circulating pag. 11 in the plasma and has to be maintained at low concentration (0,1 mg/dm 3) because it is toxic at high concentration (>150 mg/dm3). The matu re red blood cells have no nucleus . The biconcave shape also provides a greater surface area across which gas exchange can occur, relative to its volume; a sphere of a similar diameter would have a lower surface area -to-volume ratio. Their cellular membran e is called stroma and it has a variable thickness: it is thicker in the periphery and thinner in the biconcave portion. Shear stress on RBC Blood viscosity In order to calculate the blood viscosity, we have to start from the water viscosity. According to Poiseuille, the water viscosity can be obtained as: pag. 12 The vessel diameter is another variable that affects whole blood viscosity. Indeed, as the diameter of the capillary decreases, for example in capillaries, the size of the corpuscular elements (erythrocytes, leukocytes...) can no longer be neglected from the point of view of interaction with fluid -dynamics. Coagulation Each time blood pass across a not biological surface it starts to coagulate: coagulation is an ir reversible process that should be avoided in the external circuit of the extracorporeal circulation, in fact only a small part of the coagulation plaque can be re -converted into blood. The process is sometimes characterized as a cascade, because one event prompts the next as in a multi -level waterfall. The result is the production of a gelatinous but robust clot made up of a mesh of fibrin —an insoluble filamentous protein derived from fibrinogen, the least abundant plasma protein produced by the liver —in w hich platelets and blood cells are trapped. Coagulation can also take place inside the physiological system as part of the hemostasis, for example in case of the atrial fibrillation (platelets promote coagulation). Atrial fibrillation must be avoided becau se its effect is that the blood is not pumped outside the chamber (poor pumping action, trembling of the heart walls, the blood is too much static inside the heart chambers, in this way the intervention of platelets starts the coagulation). If a clot is fo rmed in the heart ventricle, it will be pumped in the circulation, causing strokes. Blood conducting structures must have a fluid dynamic design in order to avoid vortexes (because in case of vortexes we have 0 velocity in the internal points and stagnatio n is shown, in this case coagulation can take place), abrupt variations of section or path that can determine vortexes and static points. The shape has a fundamental function and if a change of shape is needed it must be gradual. If the heart function is p roviding a non -correct circulation, anti -coagulants and antiaggregant are needed (otherwise the blood clothing formation is present). Coagulation is depending on the fluid dynamics of the circuits, the devices must be designed in the proper way, to minimiz e this kind of phenomenon. To avoid this kind of process the typical drug that it is used is heparin (now it is a synthesized drug but once it was derived by the liver of the horses). The patient after its administration is resulting as hemophilic. There pag. 13 are two kinds of heparin: heparin calcium and heparin sodium, the latter is injected in the blood. Heparin is used not only during extra corporeal circulation but also in resting condition, when the patient must stay at rest (injected and absorbed). This problem increases when we use not biological materials, in this case the blood starts to coagulate because it doesn't recognize as self any artificial material. The blood covers all the foreign surfaces by platelets: in this way the surfac e of the material reproduces the neointima layer so that the blood can recognize the material as biological. This process is not controllable and each time blood operates in this way, there is an extra -coagulation that tends to close the vessel or reduce t he caliber of the vessel. To avoid an enormous coagulation, vascular prostheses are subjected to an operation called re -clotting (in case of Dacron tubes): the prosthesis is kept in contact with the blood of the patient for a certain time so that blood pop ulates the material and promotes the coagulation and form a new vascular endothelium (self -destroying aspect of the body, that doesn't recognize the help of the new device that is implanted and because of that on this kind of implants we observe coagulatio n). After the surgical intervention, there is the need to revert the action of heparin in the circulatory system, another drug is needed to do this kind of action is the protamine (it is not possible to wait until the system has metabolized the heparin): the system come back to normal coagulation but it is quite toxic, that's why the patient is administered with small quantities of heparin over time so that it is enough to administrate to the patient low dose of protamine (otherwise anaphylactic shock will be present). There is no known material, natural or ar tificial, that does NOT cause blood coagula tion. There are only materials that make it coagulate more slowly than others. The only current solu tion is to interrupt the coagula tion chain, for example by administering heparin, which blocks the transforma tion of fibrinogen into fibrin. Coagulation and biomaterials Suppose we have an arterial blood vessel isolatedfrom the body. If you put blood inside it, does it clot? The answer is yes! Why? - The vessel dies : it is no longer nourished by vasa vasorum; the material, even if natural, no longer communicates with the blood , so t he coagulation chain is triggered, although more slowly than in other materials - Alteredfluid -dynamics Fluid -dynamics and coagulation A non -physiological fluid -dynamic behaviour, such as a vortex, is enough to induce coagulation. The times are quite long, months or even years, but sooner or later the blood coagulates (e.g. aneurysm). An pag. 14 experiment done by Vorhauerand Taray at the beginning of the 1970s showed that there is a correlation between fluid dynamics and coagulation. Several objects of the same material but with different shapes characterized by all having the same frontal cross sec;on (*). These objects were weighed, then placed in the aorta of dogs and left in -situ for a specified period of time. After that, they were extracted and weighed again. The difference in weight between before and after insertion indicated clot generation . It is observed that the order of thrombogenicity of these objects is the same as the order of the tail vor+ces, called Von Kármán vortex street. This experiment does not yet allow a quantitative correlation between the fluid -dynamic and the blood coagulation, but it does show that increased vortex formation triggers increased clot formation. It is fundamental to design cardiovascular devices (e.g. mechanical valves) . pag. 15 Haemolisys The hemolysis is the rupture of the membrane of the red blood cells (due to the elderly of t he red blood cells). There is a physiological hemolysis of the red blood cells: they live 120 days and then they die and they undergo to physiological hemolysis. The stroma rupture by itself and the substances contained in the previous red blood cells are released in the plasma, they are filtered by the liver and the kidneys and then they can be destroyed and metabolized by the spleen. Induced and non -physiological hemolysis must be avoided. If we underwent to massive hemolysis we will have a large amount o f hemoglobin moving and it will be toxic for the organs but at the moment we are not able to impede it, furthermore we will observe problem in the O2 diffusion for the lack of RBCs and a large number of stromae that will tend to occlude the kidney’s vessel s. To determine the severity of haemolysis, the amount of free haemoglobin in the blood is measured using an instrument called a spectrophotometer. Since haemoglobin is red in colour, the spectrophotometer can determine the amount of haemolysis by measuring the red intensity of the plasma. As the intensity of red increases, the content of free haemoglobin in the blood increases and, therefore, the haemolysis. The main agents that are responsible to the hemolysis are: 1. chemical hemolysis: it is cause d by bacteria, toxins and poisons (e.g. from mushrooms, snakes, drug overdoses); in this case there is a massive hemolysis and an high concentration of free hemoglobin (increasing toxicity), there can also be kidney failure because of stromae accumulation 2. Thermal hemolysis: each kind of cell decrease its properties in case of the rise in temperature until 38°C and at 42°C the cells die; not only the temperature but also its gradient is also important (defined as ������������ /������������ ), that describes how rapidly temper ature changes over time and it is the main responsible for thermal hemolysis. If the increase of temperature is too fast, the red blood cells will be subjected to hemolysis, but this won’t happen in case of a rapid decrease in temperature: red blood cells are more resistive to cooling than to warming. Re -warming of the blood following different steps requires low gradient (< 3,4°C/5min), otherwise coagulation and hemolysis will be present. Cooling can be done quite rapidly. The temperature difference betwee n the cooling/warming phase and the blood mustn’t be higher than 10°C. During extra corporeal circulation heat exchangers are needed: the patient is opened in the chest and the internal temperature of the patient falls down at 34 °C just because of the tem perature inside the surgical room that is around 18 -20°C); furthermore if we need to operate close to the heart in some cases (for example in case of aneurism of the ascending aorta), the reduction of temperature reduces the metabolism, in any case there i s the need to re -warm the blood (extracorporeal circulation is used in case of people that have to be re -warmed after freezing condition, low temperature allow to survive) 3. Mechanical hemolysis: caused by the use of non -isotonic (for example water) solution s put in contact with blood, it happens because the liquid inside the red blood cells have a high concentration of sol. In case of administration of water, the solution results as diluted because there is an high concentration of sol inside the red blood c ell and a low concentration outside the cells: the water will start flowing inside the red blood in order to reach the equilibrium condition and we will have the rupture of the red blood cells (hemolysis due to osmotic effect, it has a swallowing effect an d causes a variation in shape of the red blood cells, they will rupture in the portions where the thickness is lower) pag. 16 Tillman diagram Tillmann Diagram puts together the evidence of haemolysis coming from several researchers. Tillman finds that the rupture of RBCs stroma is proportional to: - Magnitude of the stress - Duration of the stress The following formula links shear stresses to time: The previous equation, in a double logarithmic diagram, represents a line that bisects the plane stress -time. Instead of a hard line, a range of values should be considered, since the average lifetime of an erythrocyte is being referenced. The shear stress can be estimated knowing the flow rate flowing in the device and its geometry. In cylindrical vessels/tubes we have already seen how to calculate the maximum stress, in more complex cases it will be necessary to implement CFD modelling studies. The Tillmann diagram allows to estimate: - The induced haemolytic damage, knowing the value of the stress and its application time. - The application limit time to have acceptable values of haemolysis. - The maximum acceptable stress knowing the duration of the stress. Main disadvantages : - Variation in the threshold: the animals could have been frightened or had just eaten or were fasting - High variability (other studies demonstrate the influence of these factors on erythrocyte resistance). pag. 17 CARDIAC PHYSIOLOGY Anatomy of the heart Venous blood returns to the right atrium (RA) via the superior (SVC) and inferior vena cava (IVC). Blood passes from the RA into the right ventricle (RV), which ejects the blood into the pulmonary artery (PA). After passing through the lungs, the blood flows into the left atrium (LA) and then fills the left ventricle (LV), which ejects the blood into the aorta (Ao) for distribution to the different organs of the body. The image shows the major coronary artery . In this figure, approximate conduction velocitie s of different regions are noted in parentheses. Note that Purkinje fibers have the highest conduction velocity and the atrioventricular (AV) node has the lowest conduction velocity. pag. 18 Parallel arrangement of organs within the body. One major exception is the hepatic (liver) circulation, which receives blood flow from the hepatic portal veins of the gastrointestinal (GI) circulation (series) and from the aorta via the hepatic artery (parallel). The other exception is the bronchial circulation (from the aorta to the pulmonary veins –not shown) . Cardiac output (CO) The primary function of the heart is to impart energy to blood to generate and sustain an arterial blood pressure sufficient to adequately perfuse organs. Each time the left ventricle contracts, a volume of blood is ejected into the aorta. This SV, multiplied by the number of beats per minute (heartrate, HR), equals the cardiac output (CO) . Effect of preload on stroke volume Preload is the initial stretching of the cardiac myocytes prior to contraction; therefore, it is related to the sarcomere length at the end of diastole. When the mechanical properties of isolated cardiac muscle are studied in the laboratory, we find that if the muscle is stimulated to contract at low resting sarcomere lengths (i.e., at low preloads) under isometric conditions (fixed length), the amount of active tension developed (total tension minus the resting tension) is relatively small. If the same experiment is repeated with the muscle at a longer pre load length prior to contraction, the developed active tension is increased. If this experiment is done at several different preload lengths, and active tension is plotted as a function of preload, we observe the relationship shown in the figure. This plot is called the length -tension diagram for isometric contractions. pag. 19 There is no single, unique active tension curve in the length -tension relationship. The slope and magnitude of the active tension curve at a given preload depends upon the inotropic state of the muscle. If, for example, inotropy is increased by applying norepinephrine, the total tension curve shifts up and to the left as shown in the figure. The length –tension relationship, can be applied to the whole heart. Sarcomere length cannot be determined in the intact heart, so indirect indices of preload, such as ventricular EDV or pressure, must be used. Frank -Starling mechanism When venous return to the heart is increased, ventricular filling increase, and therefore its preload. This stretching of the myocytes causes an increase in force generation, which enables the heart to eject the additional venous return and thereby increase SV. This is called Frank -Starling mechanism in honour of the scientific contributions of Otto Frank (late 19th century) and Ernest Starling (early 20th century). In summary, the Frank -Starling mechanism states that increasing venous return and ventricular preload leads to an increase in SV. Therefore, the Frank -Starling mechanism plays an important role in balancing the output of the two ventricles. pag. 20 Guyton model The graphs below show the cardiac pressure and the cardiac volume tracing in the Wiggers diagram. The heart is characterized as a pump not only by the energy point of view but also in its capacity to self - adjust, and this behaviour is studied by the so -called Guyton Model. Guyton studied how the cardiac output (CO) is regulated as a function of the system that is perfused by the flow rate produced by the heart it self. The Guyton diagram was obtained experimentally, and the main aims was to determine how the cardiac output vary, as a function of the heart rate (beating), the afterload, and the possible coronary stenosis. pag. 21 Guyton want study the relationship between CO and VR, specifically : To this purpose, he use and animal Heart -Lung complex . As a preload he use a very large reservoir mimicking and adjusting the Right Atrium Mean Pressure (RAMP) by varying the height of it. At the outlet of left ventricle he put a variable clamp to change the afterload. In addition he use a pacemaker, to change and stabilized the heart rate. CO -RAMP relationship Cardiac out put is plotted as a function of Right Atrial Mean Pressure (RAMP); normal (solidblack), enhanced (red), and depressed curves (red) are shown. Cardiac performance, measured as cardiac output, is enhanced by an increase in heart rate and inotropy and a decrease in afterload. pag. 22 VR -RAMP relationship The venous return is given by: In this case the Mean Systemic Pressure (MSP) has to be considered: As volemy increases, MSP will increase, and consequently the venous return. MSP is the up stream pressure of the venous return, it is a reference pressure. It can be considered as the pressure that one global organ (corresponding to all the organs in the body) would have. CO -RAMP & VR -RAMP relationship In normal and physiological conditions the behavior of the heart followed blue line. At the end of the curve, a decrease in the CO is shown because an increase in the RAMP means that the heart is less fed with oxygen and the muscle contracts not in a correct way. An increase in the RAMP (higher pressure to win, for example in c ase of coronary stenosis) means that there is not the correct energy to be given to the flow rate. Increasing the frequency of stimulation of the pacemaker, the cardiac output diminishes, due to the relationship between the cardiac output and the stroke vo lume . pag. 23 During Extra Corporeal Circulation (ECC) the patient is connected to the extracorporeal circuit, creating a new circuit. In this case the total volemy(Vtot) is given by : pag. 24 HEART VALVE PROSTHESIS Which valves need more frequently to be substituted? Heart valve disease The main heart valve diseases are: - Bicuspid aortic valve o Two aortic valve leaflets instead of three o Congenital (from birth) - Rheumatic heart disease from rheumatic fever (when strep throat and scarlet fever infections ar e not treated properly) - Age –sclerosis (thickening from aging) - Endocarditis (infection of the heart valve) , they can cause stenosis or regurgitation Criteria design of a HVP The criteria are: - Unidirectional Valve - Passive (as natural valves are) - “Light” -> low density - Biocompatible (to enhance the interaction with the blood) - No Mechanical Failure pag. 25 - Good Fluid -Dynamics (to reduce clots formation) - Embolism History Complication that may occur with mechanical or tissue valve implantation : - Blood c lotting - Stroke (clots that migrate) - TIA (temporary stroke) - Clot on the valve (thrombosis) - Bleeding - Haemolysis - Infection - Scar tissue growth - Reoperation - Tissue degeneration (tissue valve only) - Mechanical Failure - High pressure drops pag. 26 Caged -ball The first artificial heart valve was the caged -ball (Starr -Edwards), which utilizes a metal cage to house a silicone elastomer ball. When blood pressure in the chamber of the heart exceeds that of the pressure on the outside of the chamber the ball is push ed against the cage and allows blood to flow. At the completion of the heart's contraction, the pressure inside the chamber drops and is lower than beyond the valve, so the ball moves back against the base of the valve forming a seal. The ball tends to bum p before coming to the complete opening and when this happens there is a variation in the cross section in which the blood can go through, causing an increase in the shear stresses (if we vary the cross section, the velocity changes and we will increase th e risk of hemolysis). When reducing the section, there will be an increasing velocity (the flow rate remains the same) and so increasing shear stresses. The ball is kept in position by the lower cage, but also during the closure there will be some rebounds and so hemolysis. No passage of blood is allowed during the closure of the valve, the cage have to be high enough to guarantee the flow of the blood and making the valve competent (in case of low profile gate, the valve would remain closed and the passage of blood is impeded and so low profile is not possible). These kinds of valves guarantee unidirectional flow but the main problem is that the silicon ball may reduce its dimension after some years because of the contact with blood, and it can esca pe from the cage (in this case the valves can be substituted only if the valve was in the mitral position, in the aortic position the patient would die immediately). The second prosthetic valve is called Smeloff -Suttor and it is made by two cages (upper an d lower) and they are cut in the end. The valve protrudes in the chambers very much (incongruent in the outlet valves, less incongruent in the inlet valves and so in the upstream chambers: the cage may touch the internal wall of the aorta making them react by producing denaturation of the components of the intima of the artery). pag. 27 Prosthetic Heart Valve HistoryFluid -Dynamic behaviour of the caged -ball valve : Tilting disc The ring contains a radiopaque agent, in order to make the prosthesis visible through X -rays for the evaluation of the correct position. The physiological valves are kept in position by the papillary muscles and by the tendinous chordae; during surgical interventions, the chordae are c ut from the leaflet but they remain attached to the papillary muscles and they can remain trapped in the cage and they can impair the closure of the valve. Like the Starr -Edwards, they are competent valves. All the mechanical valves are characterized by a mass and during closure all this kind of structures they are made of will exchange energy with the blood, furthermore there is a backflow in the closure due to the fact that they do not close immediately (because of their weight). In order to obtain a low -profile valve, a new kind of valve was produced, but at these times it wasn't possible to connect the moving part with the rigid one (no hinges). The ball was replaced by a translating disk. The cross section is large but also in this case some chords can remain trapped in the cage and the disk –made of silicon just like the previous ball - change the dimension and can escape from the cage. pag. 28 Then the tilting disk valve was introduced: in this case there is no need of keeping the disk in position through a cage, because the disk is connected with the annulus through hinges. It wasn't easy at the beginning to obtain this kind of valves, because it was difficult to keep in position the disk. The disk can move but not too much, the materials where more resistant and durable (pyrolytic carbon) and the surfaces were smooth. The tilting disk valve has a precise and correct positioning with respect to the flow (they can’t be inserted in any position). The opening an gle for the flowing of blood is not 90 degrees, even if it would guarantee the maximal cross -sectional area possible. Reaching this kind of configuration at the opening is not possible, otherwise no action of the pressure forces is obtained on the disk (th e flow is going downward, when the downstream pressure is greater than the upstream one, there is a force that insists on the valve and it needs a cross -sectional area to act on): it is mandatory that the disk doesn't open completely and because of that an impaired fluid dynamics is shown (the blood flow rate is not perfectly subdivided in the two sections). If the valve is completely open, its orthogonal section to the flow is not big enough to guarantee the closure of the valve. The surface of the disk is quite large and the inertia is high (all the components have to be put in motion and the disk is closing against a column of water). These valves are affected by a high backflow during closure because of the buoyancy effect of the disk. The disk can be fl at (not proper for a total perfusion) or curved (it tends to form vortexes in the streamilines). Bileaflet The modern mechanical heart valves are the so -called bi -leaflet valves: they allow an open section to blood. The disk is divided into two parts and the forces needed for the closure are less intense in order to cause the opening of the leaflets: they close faster and the back -flow during closure is smaller than the one of the tilting disk. These valves were designed after the possibility of manufactu ring small hinges with low friction (otherwise they will rupture). These hinges have to maintain a working capacity for long time and for millions of cycles. pag. 29 Biological Heart Valve History Biological valves are divided in three families: - Autogra ft ( From the same individual, e.g. pulmonary to aortic ) - Homograft ( From a different individual, e.g. cadaveric ) - Hetero/Xeno -graft Autograft ( From a different species, e.g. porcine, bovine ) pag. 30 Determination of the optimal diameter of a prosthetic heart valve Heart valves can be roughly considered as local geometry variations causing energy losses in the cardiovascular system. Thus, it is possible to calculate the instantaneous pressure drop through the valve by using the general equatio n: Assuming the orifice of both natural and prosthetic heart valves to be circular, we can write the velocity through the valve as: And substituting we get: If we put: We can write: pag. 31 Where k is a constant which is charateristic of the specific heart valve prosthesis model. ������p is the instantaneous pressure drop, q is the instantaneous flow rate through a valve of internal diameter Di. The average pressure drop can be calculated as follows: And where dt is the period of forward flow passage through the valve. During each phase of the cardiac cycle, the flow is a function of time. To solve equation (6) this function must be known. Swanson and Clark (1977) reported the following estimate of the relationship between the flow through the aortic valve and time during the ejection phase . Analogously for the mitral flow during the lef tventricular filling phase it is found on the basis of the flow curve reported by Ta lukder and Reul (1978) : Equations (7) and (8) give values of the left ventricle outflow and inflow, thus Qa>0, and Qm body surface area). The cannula must allow the passage of a given flow rate without inducing excessive haemolysis. Among all the components of the extracorporeal circuit, the cannula is the most haemolytic element. It has a pressure drop of 100 mmHg. To reduce the pressure drops, the cannula should have the largest diameter possib le since: pag. 40 Where Q is the Cardiac Output, v the velocity and A the cross section area. Increasing the cross section area, lower will be the velocity and the stress. But d uring weaning from the ECC, the heart , ”stressed " by the surgeon, starts beating again. If the cannula had a diameter equal to aorta one, it would encounter a very high resistance (afterload), to the point of not being able to restart. This fenomenon is known as Restart barrier. We are therefore faced with two completely confl icting needs. Possible solutions Useful cannula configurations are needed to minimize related problems , so t he surgeon should strive to: - Insert the cannula as low as possible into the blood vessel -> consequent problem of well fix the cannula, because it must be ensured that the cannula does not slip out during the intervention. - Direct the cannula into the vessel in an appropriate way. Moreover, d uring the design phase, it is necessary to give the right configuration to the exit section of the arterial cannula . The blood coming out of the same does not go to impact against the wall in the opposite direction of the insertion (impact haemolysis). Therefore, it is possible makes cannulae with the tip cut into a flute beak, like this: - The outflow section is increased - The flow is correctly directed The candidate materials are metal and plastic. Plastic Metal Pro - Deformable, which is very important from the surgeon's point of view, as the circuitry has to give as little discomfort as possible. - More resistant to mechanical stress - Lower wall thicknesses can be achieved compared to plastic cannula. Cons - It is not possible to decrease the wall thickness much. - High costs For economic reasons, all operations on adult patients are performed with plastic cannulae, while for paediatric patients, where the problem of size is crucial, metal cannulae are used. pag. 41 A so -called "tobacco purse" is created on the aorta, which is necessary to tighten the flaps of the artery on the cannula after it has been threaded. Air entry must be avoided: - On the cannula, branch for air elimination - Insertion with clamped cannula → tube and cannula facing each other → connected when blood overflows. - Embolism very risky, especially if cerebral. An alternative connection is Femoral artery: - Indications: o Aorta surgery o Surgery in patients with adhesions (already operated) - Problems: o Blood travels through the vessel in reverse direction in the descending aorta → Renal Problems, as renal arteries face down ward → lower flow rate. Venous cannulae They generally present fewer problems than the arterial ones. Regarding size, the haemolysis is the most important parameter to considerer. Having a large section is important for venous return, because a large diameter decreases pressure drop. Because the pressure in the vena ecavae is very low (~2 mmHg), it is important to minimize pressure drop to ensure proper venous return. The relative size between cannula and veins must be such that blood can enter the right atrium during the restart transient. Caval cannulas have a diameter around 2/3 the diameter of the vena cava. Caval cannulae are usually made of plastic. With this type of material it is necessary to pay attention that the tube does not "kneel", that it does not bend to the point of causing restriction that no longer allow the passage of fluid (kinking effect). Plastic tubes reinforced with a metal helix placed inside the thickness of the cannula are used. pag. 42 Connections: - Surgery outside the heart → cannulation into right atrium (auricle or appendage) - Surgery inside the heart → cannulation of the venae cavae Air removal If there are devices from which it is difficult to remove bubbles, an alternative method is to fill circuit with CO2, which is more soluble in both water and blood and then proceed to fill with priming fluid. It is more important remove the air in the arterial cannula. Reservoir The blood free level of the reservoir must be controlled in order to avoid its sudden emptying (variation in height in a cylindrical reservoir are much more controllable, because the height of the reservoir is high). Small volumes inside the reservoir are needed because they are easier to control. The gravitational venous return is given by the difference in height between the right atrium (of the patient) and the free level of the reservoir: the venous return can be changed just by changing the resistances of the tubes. Norma lly the venous return line is pre -clamped (reducing the crosssectional area of the tube -localized pressure drop -), causing an increase of the resistance: to increase the venous return, the clamp must be released (increase of the cross -sectional area). Thi s can be obtained by a clamp present in the stent or by a sort of throttle (it is a qualitative maneuver because the change in the flow rate is not controlled). If the venous return flow rate is lower than the pump flow rate, the level of the reservoir wou ld decrease; if the venous return flow rate is greater than the pump flow rate, the level of the reservoir would increase. Inside the reservoir there is a depth filter and it is needed because the blood has to be claned: in the reservoir, in fact, the veno us return is contained and sometimes it is added with small flow rates deriving from the aspirators and the suckers. The suckers are cannulae used to empty the operatory field from residual blood, leakage blood (produced by cuts) and they are necessary to collect the return from the bronchial circulation (that bypasses the cardiopulmonary bypass), otherwise the heart would be filled with blood and it will start to pump blood again. Anyway, suckers are not the ideal devices because they produce high hemolysi s: they develop a high vacuum force and the blood is mixed with air that will produce hemolysis and protein denaturation (because the flowrate is really low and it is difficult to fill the tubes). The presence of the filter in the reservoir is needed to cl ean the blood from the stroma, the denaturized proteins and other elements that can be present in the operatory field (like pieces of bone, after the sternotomy). The filters do not increase the priming volume too much. The shape of the reservoir facilitat es the emptying process. If the reservoir isn’t present, the pump will operate a suction into the right atrium and to the heart of the patient The circuit is closed on it self (closed loop) and filled with saline solution Continuous operation for bubble elimination (filter and/or free surface reservoir). Transparent circuit for bubble detection Is tapped using clamps to remove any bubbles that may be attached to the tubes. pag. 43 would be emptied and it will collapse (as a consequence no flow), an open reservoir is needed to drain blood from the right atrium (it uncouples the RA -pump system). Filters The filters used are t he depth filter and the screen filter (also called direct interception filter) . The direct interception filter has a double action: 1. To entrap small (>20μm) and light part icles that are not entrapped by the depth filter 2. To act as a further safety element in case of the depth filter is saturated The case of the filter is made by PC (polycarbonate) , it is: - Rigid - Strong - Transparent They are not so particularly expensive compared to HE, OXY and cannulas . All these devices are disposable . Pumps The circuit is charged by many pressure drops: a pump is needed to win the resistances, giving energy to blood. Cardiopulmonary Bypass (CPB) Continuous Perfusion (CP) Pulsatile Perfusion (PP) Mostly used in cardiac surgery units for routine CPB Used in few cardiac surgery units and only dedicated to experimental studies PP would be useless, since the flow would be dampened to non -pulsatile in the peripheral circulation, by the vascular compliance Pulsatility would be maintained down to the capillary district No significant differences in the clinical outcome from the use of CP or PP Evident clinical and/or experimental ben eficial effects are documented: - Improved perfusion and performances of the organs and peripheral zone - Reduced systemic inflammatory response pag. 44 The pumps used in the extracorporeal circulation circuit are continuous pumps (especially in the pressure gradien t) and not pulsatile ones. Fumero -Parenzan Pump The Fumero -Parenzan Pump, also called PULSAMATIC, was the only pulsatile pump (late’80) developed for infant and paediatric cardiopulmonary bypass. A segment of elastic tubing is compressed by a pneumatically driven pushing plate under control of a microprocessor. Pulse rate and stroke volume can be set. The pump can be synchronized with the patient's ECG for counterpulsation heart assist. A total of 87 open -heart procedures were performed using randomly either a conventional roller pump or the Parenzan -Fumero pump (respectively 39 and 48 patients). The results show increased cooling and rewarming rate and urinary output, decreased vascular resistance, intensive care unit time and need for blood transfusion in the pulsatile group compared to the continuous perfusion group. In the pulsatile group, mortality was significantly lower (10.4% vs 25.6%) and low cardiac output syndrome was less frequent in the post -operative course. Classification A pump is a device able to move fluid by mechanical action converted from electrical/chemical/thermal (etc.) energy into hydraulic energy. The relationship of pump developed head with the pump discharge flow at constant speed in general is called the pump performance characteristic. There are several classification of the pumps : Positive displacement Rotodynamic Can theoretically produce the same flow at a given speed (rpm) no matter what the discharge pressure. Thus, positiv e-displacement pumps are constant flow machines. A positive -displacement pump must not operate against a closed valve on the discharge side of the pump, because it has no shutoff head like centrifugal pumps. A rotodynamic pump is a kinetic machine in which energy is continuously imparted to the pumped fluid by means of a rotating impeller, propeller, or rotor. Are generally divided into three classes: radial flow, axial flow and mixed flow. If a rotodynamic pump is operated with very low di fferential head, it is likely the flow will increase dramatically and so will the absorbed power. Positive displacement (volumetric) Linear -type Reciprocating -type Rotary -type - Rope - Chain - Piston - Diaphragm - Syringe - Screw - Gear - Lobe - Peristaltic pag. 45 Linear -type Reciprocating Type Rotary Type pag. 46 Roller pumps Roller pumps (see Figure 2.12) are volumetric pumps (theoretically, they are capable to give energy to win any amount of afterload, they provide a flowrate independently on the circuit resistances) and they contain a length of tubing located inside a curved raceway. This raceway is placed at the travel perimeter of rollers mounted on the ends of rotating arms, that are arranged in such a manner that one roller is compressing the tubing at all times. By compressing a segment of the blood -filled resilient tubing, blood is pushed ahead of the moving roller, thereby producing continuous blood flow. The output of the rotary pump is mainly determined by the revolutions per minute ( ������) o f the pump and the volume displaced with each revolution ( ������, which depends on the size of tubing and length of the track): pag. 47 Why do the meati open at the periphery? Internal reaction of the material -> Strong compression stress in that section -> “Safe zone” for RBCs -> Backflow (Regurgitation) -> Loss of pump volume! When describing the working conditions of the pump, we address to the internal characteristic (device itself) and the external characteristics (of the circuit downstr eam of the pump). The circuit through which the flow will be pumped is displaying (distributed and local) pressure drops. The internal characteristics of the volumetric pump are vertical: with an increasing number of revolutions (������), the flow rate inc reases. For each flow rate, the pump will be capable to pump the same flow rate whatever the hydraulic resistances are (the pump gives energy to the fluid independently on the hydraulic resistances). The pump has to provide a flow with a sufficient pressur e to reach a certain flow rate and decide pag. 48 the outlet value of the pressure: this is the intended use. The circuit displays a certain amount of pressure drops (localized and distributed). The external characteristics show a quadratic behavior and they incre ase with the hydraulic resistances ( �). If we increase � (hydraulic resistances) downstream, the pump will continue to pump and the tubes will explode and the connection between the tubes will collapse, because the energy given to blood is increasing. Δ������ can also be converted into Δ� by using the density ������. Shear stresses as a function of RPM and roller to pump wall distance Both small and large meatus can create haemolysis! Why? There is the need to minimise the haemolysis in the light safe zone and not in the second point on the right, otherwise the backflow will increase. The optimal release from the total occlusive condition is 0.15 mm for 3/8" tubes. Pro Cons Reliable You have to set the correct meatus Easy to set Easy to control Use of different material for the tube with reference to the circuit line -> connection problem Low priming volume Cheap (tube) Big console Environmentally friendly (tube) Rotodynamic pag. 49 Centrifugal pump Centrifugal pumps nare cylindrical and tiny (negligible pressure drops) and the blood enters the pump from the apex (central portion) then it flows tangentially between two subsequent blades (central inlet, tangential outlet), the pump rotates and the blood is perfused in the si de branches. It normally exploits the effect exerted by diffusors (while the roller pump behaviour depends on volume displacement, because the roller changes the volume of the tubes squeezing them). We can define which is the smoothest velocity of the fluid to cross the channel (and so the most tangential way to the blade to move through the pump). The blades and the geometry of the cones are defined in a way such that they can be used to generate a flowrate. In this kind of pump, the ai ms are to minimize the pressure drops and any kind of energy loss (no vortexes and no mixing flow rate) and so laminar conditions of the flow even in the pump are imposed. Increasing the flowrate that is entering the pump the slope of the total velocity will change , no more tangential . The rotation velocity can be changed in order to obtain higher or lower flowrate. The main difference between the roller pump and the centrifugal pump is that the flowrate in the last case isn’t constant and the internal cha racteristics vary with the number of revolutions. Changing the flowrate means changing the efficiency of the pump, which is the witness of the transformation of the kinetic energy into pressure and it wouldn’t be equal to 1. A reduction in the efficiency (velocity has not been transformed so well into pressure) would produce hemolysis, vortexes and the detachment of the fluid from the walls of the channels (the fluid to come in contact with the channels later). These pumps are designed for not s o large range of flow rate. There is a short transient time in which the heart and the extracorporeal pump are working in parallel. They are thermodynamic devices because a thermodynamic transformation takes place inside the pump: the inlet velocity of th e fluid is transformed into pressure when it flows into the pump. Tiny backflows are allowed. They are open chambers pumps (while the roller pumps are closed devices) and pressure is given to the fluid because of a divergent geometry. pag. 50 We have a transfor mation from velocity to pressure. What is lost in velocity is gain in pressure. A well -designed pump is less haemolytic than a roller -type pump. But what is meant by "well designed"? It is necessary to introduce the concept of velocity triangle. Let us co nsider two reference systems, one (A) rotating around an axis and the other (B) stationary. Imagine, for example, two men, one (A) on a rotating platform and the other (B) stationary, observing the first from above. (A) travels along a straight track towards the center of the platform, but since the platform is moving, (B) sees (A) travelling along a spiral trajectory obtained by the sum of a rotational and a linear motion. Velocity triangles: Blood axial velocity V is the sum of the relative velocity W (of the blood, referred to the rotor velocity) and of the tangential velocity U. pag. 51 - Red : radial velocity, velocity of the fluid - Green : tangential velocity, velocity of the blades - Blue : absolute velocity, vector sum of the other velocities pag. 52 Oxygenator The historical limitation in the development of ECC was primarily the difficulty in designing an oxygenating system. First attempts: mimic what really happens in the human body, starting with anatomy. The lungs have a modular structure, whose elementary exchange unitis the alveolus. Blood comes into contact with the alveolar atmosphere through a small partition consisting of several biological membranes. In the organism, such an exchange mechanism requires an area of about 50 to 70 m2. Fortunately, it is possible to reduce this surface area by noting that: - The lungs are sized to provide exchange even during high physical activity, whereas the oxygenator must provide the very limited exchange of a patient in the operating room. - Gas composition is a parameter that can be set directly in the operating room, so an oxygen -rich mixture is used. Henry’s Law But, what is the Partial Pressure? In a mixture of gases, each constituent gas has a partial pressure which is the notional pressure of that constituent gas if it alone occupied the entire volume of the original mixture at the same temperature. Air Inhaled Air Average composition Gas % Pressure [mmHg] Oxygen 21 160 Carbon dioxide 0,04 0 Nitrogen 78 593 Argon 0,93 / Other 0,03 / M. Air Oxygen 20,4 -21,4 / Nitrogen 78,6 -79,6 / Alveolar atmosphere Gas % Pressure [mmHg] Oxygen 13,2 100 Carbon dioxide 5,3 40 Nitrogen 75,4 573 H2O 6,1 47 pag. 53 Gas partial pressure in the blood . These percentage values correspond to the following partial pressures for arterial blood: - pO 2 = 100 mmHg (99% saturation) - pCO 2 = 40 mmHg Partial pressures for venous blood average the following : - pO 2 = 40 mmHg (75% saturation) - pCO 2 = 46 mmHg The “driving forces ”, i.e. ������p, are equal to 60 mmHg for O2 and 6 mmHg for CO 2. To guarantee correct gas exchange a surface area of about 60/70 m2 would be necessary to oxygenate and wash out carbon dioxide. Blood can transport oxygen in 2 forms: - Dissolved in plasma - Chemically bound to haemoglobin (Hb) Fick’s Law The flow of gas that is transferred from compartment 1 to 2 is directly proportional to the difference in partial pressure of that gas between the two compartments and is inversely proportional to a resistance to flow. [Analogy with Ohm’s Law] Mass transfer Goi