Cardiovascular surgery deals with congenital heart defect repair, valve replacements, heart transplantation, pacemaker placement, normal rhythm and/or conduction restoration (arrhythmia surgery, i.e., Maze procedure for atrial fibrillation), and coronary revascularization.

The ductus arteriosus is a vessel that connects the pulmonary artery and aorta in the fetus to shunt immature lungs, gas exchanges occurring in the placenta between fetal and maternal bloods. Shortly after birth, when the neonatal lungs carry out blood oxygenation, the ductus arteriosus closes. However, in about 1 in 2000 nonpremature infants, the ductus arteriosus remains open (patent ductus arteriosus), causing vascular diseases.

Many other types of heart malformations or congenital heart diseases exist, such as ventricular and atrial septal defects, atrioventricular canal, arterial stenosis, coarctation of the aorta, vessel transposition, cardiac valve atresia, hypoplastic left heart syndrome, tetralogy of Fallot, among others. Most defects are effectively treated by surgical and heart catheterization procedures.

Cardiac pacemakers are implanted into damaged hearts to correct defects in conduction of the action potential. Sleep apnea is current in adults with pacemakers. Nearly two thirds of patients with implanted pacemakers have obstructive sleep hypopneas and apneas. Patients with pacemakers thus should be routinely checked, because untreated sleep apnea contributes to cardiovascular deterioration [757].

Cardiac transplantation is the choice therapy for evolutive heart failure.Astrategy to treat severe heart failure targets cell growth and tissue regeneration or replacement, which are in a relatively early stage of development. As the donor pool for cardiac transplantation is insufficient, the alternative strategy relies on implantation of left ventricular assist devices (LVADs).

In the presence of arterial stenoses, understanding of atherosclerosis evolution, diagnosis at relatively late stage of development, occurrence of complications in patients with moderate stenoses and not only severe ones, as well as the respective role of normal and tangential hemodynamic stresses are aspects that stimulate investigations by biomechanicians and biomathematicians involved in the implementation of atherosclerosis models as well as computer-aided medicine tools.

Late-stage intimal hyperplasia develops after grafting at the distal anastomosis and can cause restenosis, that is, renarrowing of the arterial lumen (Vol. 8, Chap. 8. Atherosclerosis—Medical Aspects). Many biomechanical studies were aimed at optimizing the neovasculature design to prevent this complication.

Antiarrhythmic drugs are only partly effective in treating atrial fibrillation. Moreover, they have side effects such as arrhythmia. Paroxysmal atrial fibrillation triggered by focal drivers usually situated in a pulmonary vein can be cured by a catheter-based ablation.

Tachyarrhythmias can be treated by radiofrequency ablation (RFA). A selected small patch ( 5 mm) of nodal cells of the conduction paths responsible for the abnormal cardiac rhythm is destroyed by a radiofrequency wave using an imageguided catheter-based procedure.

Cryoablation can also be used, providing cold to freeze and destroy cells.⁹ Damage limited to the local vasculature (without repercussion on the organ perfusion) can lead to ischemic necrosis, thus contributing to treatment efficacy. At a few millimeters from the ice ball edge, the tissue temperature is warmer, thus avoiding lethal tissue damage.

Ablation is proposed to treat reentrant tachycardias due to an extrapath in or adjacent to the atrioventricular node, junctional tachycardias, atrial flutter and fibrillation, ventricular fibrillation, and Wolff–Parkinson–White syndrome.

Although the likelihood of restenosis is reduced with respect to balloon angioplasty, stents are still prone to occurrence of intimal hyperplasia and restenosis. Drugs (e.g., rapamycin and paclitaxel) are stored in stent polymeric coating and, upon implantation, are locally slowly released, further reducing the restenosis rate.

Drug-eluting stents have a dual purpose: mechanical (like BMSs) and therapeutic to prevent restenosis by preventing smooth myocyte proliferation. The first generation of drug-eluting stents (DES) were aimed at addressing the high restenosis rate after implantation of bare metal stents. The goal of following DES generation with thinner struts and biocompatible or biodegradable polymers was to tackle the late stent thrombosis.

Antiproliferative drug-eluting stents (DES) include drug (cytostatic compound) reservoirs (coating thickness 5–20 μm). The main components that determine the performance of a drug-eluting stent are its structure and rheology as well as constituent drug, polymer, and delivery mode (Tables 9.1 and 9.2). Strut embedment in the arterial wall ranges from simple contact to total inclusion, thereby mildly affecting the drug delivery in the arterial wall.

To achieve a prescribed drug release kinetics over a given therapeutic period using a given dose, the current design strategies focus on a multiphasic release from blends of biodegradable polymers. A biphasic strategy, a first phase of immediate drug release liberate an adequate dose fraction and a second extended release provides the dose fraction required to maintain an effective sustained action, once the drug pharmacokinetics and pharmacodynamics are known.

The release pattern of paclitaxel from three neat polymer matrices, polycaprolactone (PCL), poly(lactic-co-glycolic) acid (PLGA), and PLGA polyethylene glycol (PLGAPEG) was studied [775]. These matrices are representative of a broad spectrum of biodegradable hydrophobic and hydrophilic polymers [776]. In hydrophilic polymers, the internal bounds between the chains are weakened, hence facilitating surface erosion.

Although DESs reduce the restenosis rate as they prevent vascular smooth myocyte proliferation, they provoke late-stage thrombosis resulting from endothelial injury associated with delayed or absent healing.

The same cause producing the same effect, these implanted materials damage the endothelium, especially at the insertion site, and can provoke in-stent restenosis despite braking smooth myocyte proliferation as they do not support endothelium repair. Endothelium damage and removal occurring during implantation induce a loss of a local, natural controller that impedes unwanted cell proliferation and provides an interface preventing unneeded blood coagulation. Antiproliferative drugs attenuates the division rate not only of smooth myocyte, but also that of endotheliocytes.

Depot stent is a drug-eluting stent with micrometric drug containers with given spatiotemporal release kinetics and programmable drug release. Unlike DESs that have their entire surface coated, depot stent contains a set of tiny drug reservoirs, thereby reducing the contact surface between the polymer and arterial wall that may favor inflammation and thrombosis.

The polymeric matrix and drugs fill orifices (diameter/strut width ratio 0.5), either closed-end holes (depth 0.5 strut thickness) or microchannels open at both ends. In the latter case, different types of drugs can be stored in the polymer, such as an endothelial cicatrizant in the channel part close to the wall and an anticoagulant in the part adjacent to the arterial lumen. However, such reservoirs can weaken the stent structure.

Openings can evenly spread over the entire stent surface or localize either in connectors or in bars or apices of stent rings bridged by connectors. The spacing between two neighboring reservoirs equals 150 μm. The number of containers determine the drug dose that can be stored, hence fully replacing that in the stent coating (thickness 5–15 μm).

With respect to stent without orifices, blind-hole reservoirs on connectors or ring bars affect neither the radial strength nor plastic strain, but reduce the fatigue safety factor [786]. Blind-hole containers evenly spread on the entire stent or only located in ring apices alter both the radial strength and plastic strain and lower the fatigue safety factor to a greater extent than those situated on connectors or ring bars. Throughhole reservoirs evenly spread on the entire stent or on apices degrade further stent resistance to mechanical stress. Through-hole containers on connectors or ring bars that have the maximal drug storage capacity may kept sufficient mechanical integrity [786].

Although drug-eluting, metallic stents have reduced in-stent restenosis and thrombosis rates with respect to previous generations of stents, the permanent presence of a metallic foreign body within the artery impair the vasomotor function and can favor vascular inflammation (Table 9.3). Furthermore, these stents disturb vascular healing, are responsible for side-branch jailing, and preclude late lumen enlargement, noninvasive imaging, and possible surgical revascularization of stented segments.

Biodegradable (BP) and permanent (PP) polymer coating on rapamycin-eluting stents have quasi-equivalent efficacy that is greater than that of the polymer-free (PF) stent without differences in safety [791].

Biodegradable arterial (vascular) stents (BAS [BVS]), or bioresorbable scaffolds (BRS), represent the present step of conceiving, designing, optimizing, and manufacturing stents that deliver drug and disappear when they become useless (Table 9.4). They hence restore physiological strain and stress fields and mechanotransduction capability as well as obviate long-term anticoagulant therapy.

Bioresorbable vascular scaffolds ware first implanted in patients with stable or unstable angina and silent ischemia. Second-generation resorbable arterial scaffolds are now deployed in coronary arteries upon acute ST-segment elevation myocardial infarctions (STEMI) [788, 789]. Percutaneous coronary revascularization is currently the first choice treatment for patients with, as it is better than thrombolysis. Yet, stent placement in acutely thrombotic lesions often causes late stent malapposition and stent thrombosis because of thrombus sequestration between struts and the arterial wall. Multislice coronary computed tomography (MSCT) as well as invasive coronary angiography, intravascular ultrasonography (IVUS), optical coherence tomography imaging (OCT), and coronary vasomotor testing check quality of implantation and absence of complications, in addition to the follow-up.

Bioresorbable scaffolds have a greater conformability and flexibility than conventional stents, hence lowering scaffold strut malapposition as well as alterations of the endothelium, mural stress field, and arterial geometry. However, thicker struts and crossing profile (scaffold-to-artery size ratio % [ % for DES] [788]) impede stent crossing of sharp angles and its implantation in heavily calcified or tortuous arteries.

Metallic bioresorbable scaffolds have properties similar to those of conventional metallic stents (deliverability, good radial strength, and low recoil) in addition to retarded resorption. Magnesium-based alloys can be coated with by a composite two-layer film that controls both the corrosion rate of the magnesium alloy and rate of drug release.

Polymeric bioresorbable scaffolds are made of various types of polymers, each with their own chemical characteristics and resorption time [787]. Polylactic acid, or polylactide, is a thermoplastic aliphatic polyester. The degradation of polymeric devices can be optimized by combining crystalline and amorphous polymers.

A gradual resorption happens during the first 2 years after implantation, thereby restoring after a relatively long delay the arterial geometry and physiology, including vasomotion [790]. The duration of its degradation depends on the crystallization of the polymer and varies from 2 to 4 years. Magnesium is resorbed by corrosion in 2–12 months. Magnesium scaffolds are metabolized to chloride, oxide, sulfate, or phosphate salts. The by-product in the artery is hydroxyapatite that is eventually processed by macrophages [787].

A more extensive artery preparation is required before implantation of biodegradable scaffolds. Unlike drug-eluting stents, resorbable scaffolds are less amenable to postdilation because of the risk of strut fracture. Therefore, precise estimation of the vessel size is mandatory for selecting the proper scaffold size, a nontrivial task due to eventual vasoconstriction and thrombi [790].

Deposition of a uniform fibrous neointimal layer may be associated with a late lumen enlargement [790]. On the other hand, resorbable scaffolds have downsides. Early scaffold thrombosis can occur. Despite good radial strength and low recoil, acute recoil of polymeric biodegradable stents can be observed [790]. In addition, polymeric devices that have a limited stretchability can break when they are overdilated.

Due to thicker struts (150–230 μm [DES: 80–140 μm]) and larger wall surface coverage than drug-eluting stents, resorbable arterial scaffolds can entrap and squeeze more thrombi and debris between the scaffold and arterial wall, in addition to higher risk of side-branch occlusion.

Analysis of drug-eluting bioresorbable arterial scaffolds in ST-segment elevation myocardial infarction shows a mean strut malapposition equal to  %, a mean incomplete scaffold apposition area  mm, a mean intraluminal defect area to  mm, and a mean percentage malapposed struts per patient to  % [788]. In acute ST-segment elevation myocardial infarction, implantation of bioresorbable arterial scaffolds causes acute scaffold recoil in 9.7 % patients, rate of scaffold strut malapposition in 1.1 %, and small and clinically silent edge dissections in 38 % [789]. Complete bioresorption takes at least 2 years after implantation.

Aortic valves can be percutaneously replaced by valve-containing stents in the beating heart under image guidance, without strong damage of the aortic wall and obstruction of coronary ostia. First, percutaneous aortic valvuloplasty was proposed using balloon-expandable stents [792, 793]. Catheterized implantation uses a valve prosthesis that can be sutured into an expandable stent and keep its property after crimping and re-expansion. Bovine jugular venous valves are sutured inside the stent [794]. The valve-stent assembly is deployed by balloon inflation. Due to heavy periodic loading, self-expanding stents are preferred to balloon-expandable stents [795]. A pulmonary metal stent with a valve was very recently implanted using a catheter-based method in a child with a congenital heart defect.¹¹

Percutaneous aortic valve implantation can be proposed in old patients with severe aortic stenosis, particularly patients judged inoperable or at high risk for surgical aortic valve replacement [796]. However, the global failure rate equals 33% with 25% periprocedural deaths. Moreover, vascular injury at the access site using femoral or iliac arteries can cause death. Periprosthetic regurgitation occurs in at least 10% of the cases. An atrioventricular block can necessitate pacemaker implantation. An alternative approach, like the transapical implantation, overcomes these drawbacks, especially incorrect implantation of the prosthesis [797].

Any implant device may induce several biological and mechanical disturbances when device design is not appropriate and its rheology does not match the wall properties. However, the device material is selected to avoid sensitization, cytotoxicity, and carcinogenicity.

Hemolysis occurs when a part of the device is located in the flow core.

The intact and healthy endothelium possesses an antithrombotic property, hence supporting normal blood flow by separating platelets for the extracellular matrix. Moreover, endotheliocytes release nitric oxide and prostacyclin, two soluble inhibitors of platelet activation. They also secrete high concentrations of ectonucleotide pyrophosphatase–phosphodiesterase ENPP1 that rapidly hydrolyzes extracellular nucleotides such as ADP. However, a medical device, like any implanted foreign body, provides a matrix for thrombus formation, especially when endothelial cells are damaged or removed during the intravascular implantation. In particular, stenting destroys large endothelial segments. A thrombogenic response occurs after platelet deposition and clotting activation. After the initial platelet adhesion, auto- and paracrine mediators (ADP, adrenaline, thromboxane-A2, and thrombin) amplify and sustain the platelet response, as they recruit circulating platelets to form a growing hemostatic plug.

Vessel-wall damage also generates an intimal cell proliferation and quick restenosis . Growth factors can be released from platelets and fibroblasts. An inflammation due to interactions between leukocytes and the device can occur once leukocytes and complement have been activated. Reactive oxygen and nitrogen species ( superoxide and peroxynitrite) and effectors (e.g., polyribose polymerase)¹² participate in restenosis after vascular injury. In media and neointima, NOS2 is frequently overexpressed. The serum 3-nitrotyrosine-to-tyrosine ratio (protein Tyr nitration),¹³ an index of oxidative stress mainly resulting from neutrophil activation, is correlated with the narrowing of the lumen [799]. On the other hand, the NOS–NO–sGC axis attenuates intimal hyperplasia.

Finally, the medical device disturbs blood flow. Locally, the added interface rugosity caused by the implanted device can generate local flow separations. Moreover, the abrupt transition between native vessel and stented segment induces a compliance mismatch. By means of wave reflection, it can remotely affect blood flow.

Computer-aided manufacturing allows one to test easily diverse device geometries, structures, and material properties, with an important requirement: the material matching of the mechanical properties of the biological tissues. It needs to model accurately the physical processes at the interface, because the interface between the implants and the biological tissues (blood and vascular wall) is the focus of desirable and undesirable interactions. Moreover, it must be adaptable to unsteady loading induced by cardiac contractions.