5
Biomaterials for Cardiovascular Implants
5.1 Introduction
The cardiovascular system is one of the most, if not the most, important and complex apparatus of the human body. It is the organ system that maintains the homeostasis of the organism, it stabilizes the body temperature and pH, by pumping oxygen‐rich blood, nutrients, hormones, and other important substances to the cells in our body, and then taking the oxygen‐poor blood to the lungs. Besides the cardiovascular system’s importance, however, it is also affected by many diseases that can result in the loss of cardiac function.
In the last century, progresses in the cardiovascular surgery field have been remarkable. From the end of the nineteenth century until around 1950, most operations were performed by introducing fingers or instruments blindly into the heart. The most common operation was the closed mitral commissurotomy performed with the Tubbs dilator [1]. With the invention of the oxygenator and its entrance in the market, the heart–lung machine became a daily instrument in open‐heart surgeries. Congenital diseases and heart or blood vessel problems that were previously considered impossible to cure are now treated surgically a great majority of times, for example, valvular heart diseases, arrhythmias, and diseases of thoracic aorta.
In order to treat the root of the disease, it is necessary to repair or substitute the injured tissues. Nowadays, science and technology made this possible, easy, and safe. In this contest, a variety of biomaterials have been studied and used to repair and replace impaired heart issues. Those biomaterials are used in vascular surgeries but also in cardiovascular prevention and rehabilitation; those procedures include arterial bypass grafts, endarterectomies, reparation of aneurysms, stents, pacemakers, pumps, valves, and other endovascular devices [2].
The biomaterials used are divided into two main categories: natural and synthetic [3]. Natural materials are obtained from biological sources which can derive from human donors (autogenic if it comes from the patient themselves and allogeneic if it is from the same species donor but not the same individual) or animal (xenogenic), usually cow, pig, and sheep. Examples of natural materials can be pericardia, arteries, veins, and heart valves. In the case of xenogenic material studies, pig and sheep’s hearts are chosen since their heart size, anatomy, material degradation, and endogenous tissue growth are very similar to the human’s one [4], and for this reason, it can be particularly useful for the development of new devices for replacement and repair of diseased valves [4]. Instead, synthetic materials used in the cardiovascular system primarily include the commonly used metals, polymers, and ceramics; these materials are used either to construct new devices or to coat specific surfaces to resist infection, such as in the Silzone heart valve [5]. Metals such as titanium, nitinol, stainless steel, and cobalt–chromium alloys are used passively, while metals like silver and glassy carbon are used for coating [6]. Both natural and synthetic materials have their own pro and cons and their unique characteristics, but they all have to be biocompatible, meaning that the material must not change plasma proteins, cause thrombus formation, destroy cellular elements of blood, produce toxic or allergic responses, deplete electrolytes, and be affected by sterilization [7]. In addition, the material has to be durable, have good strength, and be flexible.
Since cardiovascular diseases remain the leading cause of death in the United States, resulting in nearly US$300 billion in healthcare annually [8], it is important to understand the reasons for failure and improve the design and choice of material to prevent them. Like all materials, also biomaterials used in cardiovascular systems are susceptible to various failure modes. Although typical loads are not high, as compared to orthopedic implants, for example, mechanical and fatigue failure are the most common failure methods. The life expectation of the implant is expected to be more than 10 years, so considering 90 beats per minute, it means that in 10 years, the device has to be resistant to more than 470 million cycles [9]. Other failures might be purely cardiovascular, such as thrombosis, resulting in the occlusion of small blood vessels, and hemolysis, as a reaction to the material or its degradation products.
The aim of this chapter is to show the background and the innovation of the materials that are used in the implants for the cardiovascular system.
5.2 Different Applications
5.2.1 Stents
In many diseases, a passageway channel (e.g. blood vessel, esophagus, and bronchus) that allows body fluids and other substances to flow can collapse or can be narrowed and become substantially restricted. The restriction makes the passage for the fluid flow hard and sometimes completely blocked. This happens frequently in coronary artery diseases, but also in peripheral vessels, carotid vessels, and renal arteries.
Angioplasty is the technique that has been first used. It consists in the insertion of a small catheter with a balloon in the vessel; after the catheter reaches the obstructed location, the balloon is inflated so to push the obstruction aside and then deflated leaving the artery unblocked so to allow a smooth blood flow (Figure 5.1).
Figure 5.1 Representation of balloon angioplasty: (a) before procedure, (b) primary angioplasty, and (c) submaximal angioplasty [10].
(Copyright © BMJ).
Statistical studies demonstrated that patients undergoing balloon angioplasty showed 29.8% restenosis and 1.1% reocclusion the following day [11], meaning that the vessel becomes restricted again. Today, this technique is called plain old balloon angioplasty (POBA) [12].
These limitations have been solved by percutaneous coronary intervention (PCI), known also as angioplasty with stent, a technique that has been applied for the first time by Andreas Grüntizig in 1974 and is nowadays the preferable choice to treat patients with acute myocardial infarction or occluded leg arteries [13]. With this new technology, stent patients showed 2% restenosis and no occlusion the following day [11].
The design of the stent has to ensure biocompatibility, flexibility, expandability, mechanical compatibility, visibility under X‐rays, and it needs to have strong radial strength and small contact area and be resistant to fatigue. Since the metal in the stents can lead to inflammation and subsequent restenosis, it is important for the safety and efficacy of the device to find a metal that is biologically inert [12, 14]. In order to satisfy all these properties, the first devices used for coronary stenting were the stainless steel bare metal stents (BMSs).
5.2.1.1 Balloon Expandable Stents
Stainless Steel Stents
The first generation bare metal stents (also known as BMSs) consisted of the stainless steel alloy 316L SS, a biologically inert material, which contains iron, nickel, chromium, and molybdenum. It has well‐suited mechanical properties and excellent corrosion resistance (carbon content <0.030 wt%), making it the preferred material for this application. However, 316L SS is ferromagnetic, mostly iron in composition (60–65 wt% pure FE), and has low density (as shown in Table 5.1), making it a non‐MRI compatible and poorly visible fluoroscopic material. In an attempt to improve the ability to visualize the stainless steel stents under X‐rays, radiopaque gold markers were occasionally added, which unfortunately increased the restenosis and mortality risk [15]. Another problem is related to the release of nickel that can provoke allergic reactions and may trigger local immune response and inflammatory reactions and induce intimal hyperplasia and in‐stent restenosis (ISR) [16]. It is possible to reduce the amount of nickel concentration to 4.5–9%. However, an amount of 10–14% of nickel can be advantageous in decreasing the ferromagnetic properties of the SS by stabilizing the Fe in a nonmagnetic state [17]. Initial stents made of stainless steel have thick struts (≥0.10 mm), and because of poor flexibility and high ISR rate, stents with thin struts (<0.10 mm) were developed [18]. The stainless steel was replaced by cobalt, chromium, or other elements to have thinner struts, better radio‐opacity, deliverability, and conformability and maintain the strength. However, the ISR rate is still high at around 15% after follow‐up of six months [19].
Table 5.1 Mechanical properties of the metals that are used for making stents. (Copyright Elsevier).
| Metal | Elastic modulus (GPa) | Yield strength (MPa) | Tensile strength (MPa) | Density (g/cm3) |
| 316L stainless steel (ASTM F138 and F139; annealed) | 190 | 331 | 586 | 7.9 |
| Tantalum (annealed) | 185 | 138 | 207 | 16.6 |
| Cp‐Titanium (F67; 30% cold worked) | 110 | 485 | 760 | 4.5 |
| Nitinol | 83 (Austenite phase) | 195–690 (Austenite phase) | 895 | 6.7 |
| 28–41 (Martensite phase) | 70–140 (Martensite phase) | |||
| Cobalt–chromium (ASTM F90) | 210 | 448–648 | 951–1220 | 9.2 |
| Pure iron | 211.4 | 120–150 | 180–210 | 7.87 |
| Mg alloy (WE43) | 44 | 162 | 250 | 1.84 |
Cobalt–Chromium Stents
In order to overcome the limitations of 316L SS stents, cobalt–chromium stents have been designed. The cobalt–chromium alloy has an excellent radial strength, thanks to its elastic modulus, and has a higher yield strength, tensile strength, density, and better radio‐opacity (as shown in Table 5.1). Not only the characteristics of Co–Cr make this type of stents widely used, but also the rate of thrombosis is very low (0.9%), comparable to drug‐eluting stents (1.0%) [20]. Moreover, it has been shown that cobalt–chromium stents are also suitable for large vessels, resulting in a moderate rate of restenosis and low rate of repeat revascularization [21].
Platinum–Iridium Stents
Platinum–iridium alloy was also used for making BMSs. It consists of 90% platinum and 10% iridium; it has excellent corrosion resistance and radio‐opacity (as shown in Table 5.1), and in fact, it is even possible to use MRI to take images in three dimensions of the lumen of these stents [22]. Compared to 316L SS stents, PT‐It alloy in MRI produces much lower artifacts [17], while the percentage of recoiling is higher (16%) than 316L SS (5%) [23]. Although it has poor mechanical properties, studies demonstrated that this stent was safe in animal models, with good short‐term patency without anticoagulation. From histological examination, it has been seen that on the stent site a neointimal thickness of 325–650 microns has been formed [23].
Tantalum Stents
Tantalum has been coated on 316L SS surfaces to improve corrosion properties, thanks to its highly stable surface oxide layer, enhancing the biocompatibility of 316L SS. Tantalum is known to have optimal biocompatibility [24], and its high density makes the stent to have excellent fluoroscopic visibility. However, tantalum’s yield strength is very close to its tensile strength as it can be seen in Table 5.1, increasing the possibility of breaking during the deployment. Compared to 316L SS, biocompatibility and visibility properties are greater; however, its high elastic modulus, bioinertness, and unsatisfactory osseointegration limit its wider use in clinical applications [25].
Titanium Stents
Another material known for its optimal biocompatibility and corrosion resistance is titanium. In fact, this material and its alloys have been widely used in both dental and orthopedic applications, but rarely for making stents. Titanium has a very high yield strength, comparable to the Co–Cr alloy, but with much lower tensile strength (as shown in Table 5.1), increasing dramatically the probability of tensile failure when the stent is expanded. Moreover, titanium stents have a low ductility, making the stent more prone to failure. For these reasons, it has not been developed as a purely titanium stent. However, alloying titanium with other elements that reduce its yield strength might overcome these limitations and make it mechanically acceptable while maintaining the same tensile properties [26]. Some of the alloying elements are nickel and tantalum.
5.2.1.2 Self‐Expandable Stents
Nickel–Titanium Stents
The Ni–Ti stents are a peculiar type of stents, called self‐expandable. The reason is that Ni–Ti stents do not need a balloon to inflate them, but they expand themselves due to their shape memory effect. The nickel–titanium alloy is also known as nitinol, composed of 55 wt% nickel and titanium balanced. Ni–Ti stents commonly have short struts [27], but their properties are strongly dependent on the processing history. Nitinol stents are defined as super‐elastic since they are able to recover more than 10% strain [28, 29]. If a stress is induced in the stent, the material changes shape immediately, but once the stress is removed, the original shape is recovered [30] (Figure 5.2).
Figure 5.2 Steps of the finite element model to simulate the deployment of the self‐expanding nitinol stent into the curved/kinked patient‐specific vessel: (i) stent crimping with radially displaced rigid plates; (iia) stent morphing with a cylindrical surface; (iib) stent deployment [31].
(Copyright © Creative Commons CC‐BY).
Self‐expanding nitinol stents are manufactured at a larger diameter than the target vessel [30]. Generally, they are crimped at or below room temperature, that is lower than martensite finish temperature Mf, constrained in a retractable sheath, and then implanted in the body to the target site. This operation is done to prevent a premature expansion during delivery to the body. Once it reached the treatment site, it is released from the delivery system, and due to the body temperature, typically the transformation temperature is set to 30°C or over, and the stent expands and regains its original diameter, transforming to its austenite phase [29]. The process is possible, thanks to the shape memory property: at a temperature T1 (i.e. body temperature) which exceeds the austenite finish temperature Af, the alloy is stress free and fully austenitic. When the temperature is cooled down to T2, lower than the martensite finish temperature Mf, the alloy remains stress free and transforms into martensite. During loading at T2 to C–D branch, martensite deforms due to the movement of the twin boundaries; the less favorably oriented martensite converts to more favorable oriented ones. Instead, the unloading along the D–E branch leaves a residual strain at temperature T2, which survives heating up to austenite start temperature and then completely eliminates during additional heating to T1 [32]. The process is schematized in Figure 5.3.
Figure 5.3 Shape‐memory diagram of nitinol [32].
(Copyright © Elsevier).
Nitinol is super‐elastic with optimal shape memory and biocompatibility and has great corrosion and fatigue resistance. However, corrosion can dissolve nickel and cause adverse effects [33], like allergic, toxic, and carcinogenic reactions. It has been observed that after an appropriate passivation of its surface, the Ni–Ti layers consist mainly in titanium‐oxide layer (TiO2). This highly stable and biocompatible Ti‐based oxide layer can increase the stability of the surface layer by protecting the bulk material from corrosion, but it can also create a barrier against Ni oxidation [34, 35]. In addition, corrosion resistance of nitinol stents can be further enhanced by different treatments such as electro‐polishing, which promote a very uniform oxide layer. Both methods showed good compatibility and did not promote toxic or genotoxic reactions in physiologic environments [36]. WALLSTENT is the first self‐expanding, stainless steel wire‐mesh structure, coronary stent manufactured by Schneider AG implanted in the human body in 1986 [37]. This stent had a high incidence of subacute stent thrombosis; in fact, the first coronary stent that was approved by the US Food and Drug Administration (FDA) is a balloon‐expandable BMS, developed by Palmaz‐Schatz (Cordis Corp; a Johnson and Johnson Company) [38].
5.2.1.3 Drug‐Eluting Stents
Drug‐eluting stents (DESs) are a novel and innovative approach in stent technology, introduced on the market only in 2002 [39]. The aim of the invention was to overcome the high probability of restenosis that BMSs create during PCIs [40]. In fact, the idea was to have a stent system that is able to deliver drugs locally to inhibit intimal thickening by interfering with different pathways involved in the development of inflammation, migration, proliferation, and/or secretion of extracellular matrix [41]. Both the drug and the delivery vehicle must fulfill pharmacological, pharmacokinetic, and mechanical requirements.
DES shares the same stainless‐steel scaffold of BMS, but in addition has also immunosuppressant, cytotoxic, or antiproliferative drug to inhibit neointimal hyperplasia and a polymer coating, which acts as a drug reservoir to ensure drug retention during deployment and a uniform distribution on the stent [38]. The first generation metallic stent scaffolds were made of nitinol or stainless steel and were loaded with sirolimus or paclitaxel; both of them are able to block cell cycle progression, therefore the inhibition of smooth muscle cell proliferation. These two stents had been named sirolimus‐eluting stents (SESs) and paclitaxel‐eluting stents (PESs) [37] and had become the most widely used devices to treat coronary artery disease. Sirolimus is a natural macrocyclic lactone and paclitaxel is a cytotoxic agent against many tumors. Both SES and PES were designed in stainless steel; SES (Cypher stent, by Johnson & Johnson) was coated with poly‐(n)‐butyl methacrylate, an inert synthetic permanent polymer, containing 140 μg/cm2 of sirolimus, that allows a slow release of 80% drug within 30 days after implantation, while PES (Taxus stent, by Boston Scientific Corp) was coated with a single layer of Translute, a permanent polymer coating, combined with 1 μm/mm2 paclitaxel, which allows the release of 10% of paclitaxel in the two weeks after implantation, while 90% remains in the polymer [17, 42].
Despite first generation DES demonstrated to be an optimal choice due to its high success rate and very low ISR rate [38], it revealed late (more than 30 days) and very late (more than 12 months) stent thrombosis [43], as shown in Figure 5.4. The cause of stent thrombosis is due to the antiproliferative effect of DES, which slows down the re‐endothelialization process of the prosthetic material. To prevent the formation of platelets, an oral antiplatelet therapy is required. Once it ends, the uncovered scaffold material can trigger platelet activation, and late restenosis or thrombosis occurs [45]. While in contrast, restenosis caused by neointimal hyperplasia is a slow process, with respect to stent thrombosis that occurs suddenly with acute life‐threatening symptoms. The incidence is low, but the mortality rate is high. In addition, anticoagulation might be crucial after the implantation of the stent [46].
Figure 5.4 Stent thrombosis (a) drug‐eluting stent implantation in the proximal left anterior descending artery, complicated by (b) stent thrombosis occurring seven months later, shortly after discontinuation of dual‐antiplatelet therapy. (c) Gross pathology example of stent thrombosis [44].
(Copyright © Elsevier).
In order to improve the DES performance and to achieve more desirable characteristics of flexibility, radial strength, trackability, and biocompatibility, more effort was put into the improvement of the scaffold, polymer, and drug. In this way, second‐generation DESs have been developed. Those stents were designed to overcome safety issues in the long term, employing new biocompatible polymer coatings and less toxic proliferating drugs [37, 47]. The second‐generation DESs have a thinner strut thickness, which can accelerate the healing and endothelialization of the coronary arteries and reduce inflammation and injury to the media. The new generation DESs have a cobalt–chromium platform, instead of stainless steel, improving the flexibility and deliverability of the stent. Moreover, fluorinated polymers have been used, thanks to their better biocompatibility and stronger thromboresistant properties [37].
Another factor that can induce local responses and alter processes involved in the neointimal formation is introduced by the presence of permanent polymers in vessel arterial walls. Each polymer can evoke a different inflammatory response, for example, giant cell infiltration around the stent struts or a progressive granulomatous and eosinophilic reaction [42, 48]. Therefore, various permanent (biostable) and biodegradable polymers have been used on DES platforms. Permanent polymers were the conventional polymer choice due to the ease of the drug release and the durability after drug elution is complete. These polymers have been used in first‐generation SESs and PESs, but it has been observed that those can cause delayed healing and impaired stent strut endothelialization and provoke chronic inflammation that triggers impaired arterial healing [49] and also a degradation of poly‐(n)‐butyl methacrylate over time, which can cause vascular smooth muscle cell apoptosis [48].
Second‐generation DESs have more biocompatible non‐erodible polymers, which have been shown to have greater degrees of re‐endothelialization compared to the first‐generation DESs. Polyvinylidene fluoride, hexafluoropropylene, and polyvinylpyrrolidone are the most currently biocompatible polymers [50]. However, stent thrombosis remains to be an unsolved issue in second‐generation DESs. Consequently, polymer‐free DESs have been introduced as a solution to this problem.
5.2.1.4 Biodegradable Stents
Nowadays, the metals that are mostly studied for both cardiovascular and orthopedic applications are magnesium (Mg), iron (Fe), and zinc (Zc) because they offer good in vivo biocompatibility, controlled degradation profile, and sufficient mechanical strength to support bone during the regeneration process [51]. With the introduction of biodegradable stents, problems related to late stent thrombosis, in‐stent restenosis, chronic inflammation, and stent strut fracture reduced significantly. Moreover, if biodegradable stents are used in children or infants, once it disappears, the vessel keeps growing naturally until adulthood, without requiring stent balloon dilatation or surgery in the future [52].
Magnesium Alloy Stents
Biodegradable magnesium alloy stents (BMASs) are the most studied biodegradable scaffold. Originally, magnesium has been firstly used in biodegradable orthopedic implants and have been later introduced in coronary stents. Magnesium shows values very close to the bone: high specific strength, low elastic modulus (41 GPa), and low density (1.74 g/cm3) [53]. However, pure Mg was not suitable for stents’ design requirement [54] because it has a very fast corrosion rate in physiologic environments, especially biological fluids, which leads to a huge release of magnesium and premature loss of mechanical strength of the implant [55] (Figure 5.5). Therefore, magnesium is alloyed to control its degradation kinetics. The two Mg‐based alloys used are AE21 and WE43. AE21 contains 2% aluminum, 1% rare earth metals, and 97% magnesium [54], while WE43 contains 4% yttrium, 0.6% zirconium, 3.4% rare earth metals, and the rest is magnesium [17] (Table 5.1). However, also Mg‐based alloys deteriorate over time and its corrosion rate is faster than that of bone healing rates [56]. Therefore, biodegradable magnesium alloy stents should have a much higher inherent strength to make it more time lasting [57]. Another limitation is nonuniformity corrosion, specifically pitting corrosion, that has been demonstrated both after 18 weeks, implantation in guinea pigs’ femora and within the bone [58].
Figure 5.5 Degradation of magnesium staples under in vitro and in vivo conditions. Optical visualization of morphology of Mg staples after immersion in simulated body fluid at pH of 4 for (a) 3, (b) 7, (c) 11, and (d) 14 days. Photographs of Mg staples implanted into an animal model for (e) 7 and (f) 90 days, and (g) Ti staples implanted for 90 days [51].
(Copyright © Creative Commons CC‐BY).
Iron Alloy Stents
Since iron is an essential trace element for the human body, in 2001, the idea of creating a Fe‐alloy‐based stent was studied for the first time by Trumbo et al. [59]. An iron stent had been implanted for the first time in the descending aorta of a New Zealander white rabbit. It has been observed that there are no restenosis effects and inflammatory responses in the long term. Iron and its alloys have high ductility and high radial strength; thus, it is possible to use thinner struts, reducing in this way the restenosis rate. However, Fe‐alloy stents cannot be degraded over the desired working time, leaving a large portion intact one year after implementation [60] and a large volume of iron oxide that is not safely metabolized by the body.
Zinc Alloy Stents
Over the last two decades, the research focus was on Mg and Fe stents. Fe‐based alloys showed appropriate mechanical properties but have inadequate degradation rate, while Mg‐based alloys have too high degradation rate in physiological environments and insufficient mechanical properties, and corrosion is almost never uniform. Zc alloys have been explored as a potential material for bioabsorbable vascular stents only in the last few years due to their great corrosion resistance, tunable mechanical properties, and a degradation rate of only 0.02 mm/year [61]. Despite a lower toxicity limit with respect to magnesium, the toxicity threshold is not considered as a limitation since the stent has a very small volume [62]. Moreover, zinc has a low melting point (420°) and low reactivity in the molten state, permitting melting and hot processing in air [63]. The drawbacks of this biomaterial are very similar to those of Mg‐based stents; however, research studies are still very preliminary, and extensive investigation is continuing in this field toward getting proper composition and meeting the criteria to have better biocompatibility, prolonged mechanical integrity, and controlled corrosion rate, in vascular stents [63].
5.2.1.5 Latest Stent Inventions
In recent years, researchers have moved part of their attention to develop new technologies to treat peripheral artery diseases due to the increasing number of peripheral interventional procedures required and to prevent critical limb ischemia, amputations, and high restenosis rates [64]. With respect to coronary arteries, peripheral arteries are in a much more hostile environment due to the high risk of external forces and pressures applied from outside the body, such as due to the flexion of the knee, which can break the stent and lead to high rates of in‐stent restenosis [65]. The first drug‐eluting self‐expanding peripheral stent, introduced in 2010 in Europe, is the Cook Medical Zilver PTX stent [64]. This stent is flexible, made of nitinol, and has a polymer‐free paclitaxel coating, the same one used in PESs. The results from the study of Dake et al. showed that Zilver PTX stents have excellent integrity and a good performance, making the use of this kind of stent very favorable to treat restenotic lesions of peripheral arteries [66]. A comparison has been made by IMPERIAL, a global randomized controlled trial, between the polymer‐free Zilver PTX and Eluvia drug‐eluting vascular stent, a paclitaxel‐eluting nitinol stent produced by Boston Scientific. After 24 months’ trial, the results show that Eluvia stent, which uses a drug–polymer combination, can provide a sustained release of drug for over one year [67, 68] and has better results, in terms of number of patients that had no major adverse events at 12 months, with respect to those who received the Zilver PTX [64]. However, in 2019, Dr. Dake presented a new five‐year data on Zilver PTX that confirms that Zilver PTX and paclitaxel dose were not predictors of mortality and can provide long‐term benefits [69]. The stent that has been introduced most recently is the BioMimics 3D vascular stent system; the characteristic point of this stent is that it has a helical shape and is able to reduce its length during knee flexion [64].
From coronary stents’ front, new directions are toward developing better bioresorbable and bifurcated stents. For bioresorbable stents, Abbott introduced their Xience stent, a bioabsorbable everolimus‐eluting coronary stent system, for ABSORB trial, which considers patients with single de novo coronary artery, and this one gave optimal results [70]. Xience Sierra is their latest upgrade, and this one showed an ultralow crossing profile that increases deliverability and flexibility even more [71]. Moreover, another new introduction to the market has been drug‐coated balloons (DCBs), which technology is very similar to plain balloon angioplasty, but in addition there is an antiproliferative medication coating on the balloon which helps to prevent restenosis [72].
5.2.2 Prosthetic Heart Valves
The job of the heart is to work as a muscular organ to pump blood throughout the body, so to be able to transport the oxygen and the other nutrients to every corner of the body and to carry out the waste. Electrochemical impulses created by pacemaker cells control the contractions of the heart, which drive the blood to all over our body, and ensure a rhythmic and synchronized contraction of the cardiac muscles. Thanks to the heart valves, the heart’s contracting chamber turns into pumps.
The heart has four valves, two for each side of the heart, ensuring that each muscle contracts and produces efficient unidirectional flow. In the right side of the heart, blood is returning from the body and going to the lungs to be oxygenated, crossing the tricuspid and the pulmonic valves. On the contrary, in the left side, oxygenated blood is controlled by mitral and aortic valves and is pumped in the peripheral sites [73]. The aortic and pulmonic valves open during systole when the ventricles are contracting and close during diastole when the ventricles are filling through the open mitral and tricuspid valves. A schematic is represented in Figure 5.6.
Figure 5.6 Four valves of the heart, visible with the atria and great vessels removed (credit: OpenStax College, CC BY 3.0, via Wikimedia Commons).
Heart diseases are currently the major cause of death in the world [74], and valve malfunction is one of the main causes. The most commonly affected valves are the mitral, tricuspid, and aortic valve. Those malfunctions affect the hemodynamic performance in both the forward and reverse flows: stenosis and incompetence [75]. The stenosis occurs as a narrowing of the valve, which makes the flow of the blood harder and therefore generates a larger pressure drop across the valve. However, the incompetence happens when the valve leaflets fail to close completely, causing retrograde blood flow (regurgitation) when the valve should be shunt [76]. Since these pathologies occur very frequently, prosthetic valves are very useful and widely used.
The first clinical use of a prosthetic heart valve happened in 1952 when Dr. Charles Hufnagel implanted a caged‐ball valve in the descending thoracic aorta in a patient with aortic valve disease [77]. In the last 70 years, more than 50 different cardiac valve prostheses, which differ in geometry, number of leaflets, or material, have been studied and introduced, but it is still not possible to eliminate all the problems associated with the heart valve prosthesis which include hemolysis, thromboembolisms, bleeding complications (i.e. anticoagulant‐related hemorrhage), and prosthetic valve endocarditis, followed by structural and nonstructural prosthetic valve dysfunctions [78].
Heart valve diseases are generally congenital, meaning that the baby was born with a defect to one or more valves, or originated with aging or diseases such as rheumatic fever [79]. Those defects can compromise the effectiveness of the functionality of the valve, causing the damage of the support structure or restrictions of the leaflets. As a consequence, valve stenosis or regurgitation or both eventually can cause the valve’s failure. Those effects are quantified by two parameters: the regurgitant volume, which is the flow that returns to the chamber during the closure of the valve, or also known as leakage flow phase, and the effective orifice area (EOA), which quantifies the valve opening during the forward flow phase. The EOA can be measured using the Gorlin equation,
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