Drug Eluting Stent Compounds

Drug-eluting stents (DES) are vascular prostheses used by interventional cardiologists to reopen and maintain patent coronary arteries narrowed by arteriosclerosis. The history of interventional cardiology began with balloon angioplasty in 1977. Sigwart et al. introduced the first bare metal stent (BMS) in 1986.[1] In 2002 the first DES came to the markets in Europe. Today many companies are offering diverse DES to improve the treatment of coronary artery disease.[2][3] A multitude of studies providing evidence has paralleled this development. Real world practice emphasizes the clinical importance of stent technology illustrated by the fact that stent implantation occurs in 90% of PCI.

Facing the variety of DES, one can classify them according to three characteristics:

1. Scaffold
2. Drug-delivery mechanism (i.e., polymer)
3. Therapeutic agent

The development of DES evolved through different generations. First-generation DES had a stainless steel scaffolding coated with either sirolimus or paclitaxel. The RAVEL, SIRIUS, and TAXUS trials evaluated first-generation DES.[4][5] Since studies showed the superiority of rapamycin agents, the second generation DES carried everolimus or zotarolimus. Second-generation DES has a cobalt-chromium scaffolding with different polymer coatings which allows decreased strut thickness, improved flexibility, deliverability, enhanced biocompatibility, better eluting profiles, and superior re-endothelialization. The ENDEAVOR and SPIRIT trials tested second-generation drug-eluting stents which are now the predominant implanted stents.[6][7][8] The third generation of DES with a biodegradable polymer or entirely bioabsorbable scaffolds are just undergoing clinical testing.[9]

Stents prevent the vascular recoil cardiologists observed with balloon angioplasty (PTCA). Bare metal stents (BMS) are superior to PTCA alone as was shown in the BENESTENT and STRESS trials.[10] But following BMS implantation, restenosis develops in 30% of cases. Early restenosis results from neointimal hyperplasia because of migration and proliferation of vascular smooth muscle cells as a response to vascular injury from the stent deployment.[11][12][13] To reduce restenosis rate, DES evolved from BMS. Both share the common scaffold structure. But in drug-eluting stents, an antiproliferative drug coats the scaffold to reduce cell proliferation inside the stent and also treat the complication of early restenosis.[14] This development led to a significant reduction of early restenosis. Long-term trials, however, revealed another problem, which is late (>30 days) and very late (>12 months) stent thrombosis.[15][16][17] Adequately powered long-term trials have shown that drug-eluting stents correlate with this complication. The reason for stent thrombosis is the antiproliferative effect of DES, which slows re-endothelialization of the prosthetic material. After cessation of oral antiplatelet therapy, the uncovered scaffold material can trigger platelet activation, and late restenosis or thrombosis occurs.[18][19] Whereas restenosis due to neointimal hyperplasia is a slow process, stent thrombosis occurs suddenly with acute life-threatening symptoms. It has a low incidence but high mortality.[20] Thus, anticoagulation is crucial after implantation of stents.[21] 

Problems with hypersensitivity to stents have been proposed and subsequently, the concept of polymer-free or biodegradable polymers was introduced. Recently reestablishment of healthy long-term vasomotion with biodegradable stents have gained significance. 

Different characteristics make each drug-eluting stent unique. Each of them has advantages and disadvantages which derive from differences in drug-loading capacity, drug-release pharmacokinetics, polymer durability, biocompatibility, influence on vascular wall thinning, aneurysm formation, and delayed restenosis.[22] Different complications can occur following the implantation of stents. Materials not native to the human body involve questions of biocompatibility. Problems of corrosion and release of toxic substances led to the investigation of more biocompatible and degradable compounds. Early biocompatibility problems include stent thrombosis, inflammation, and neointima formation. A late problem associated with decreased biocompatibility is scaffold fracture due to material fatigue.

The FDA MAUDE database showed that most stent fractures do not trigger symptoms, are difficult to assess and thus, are probably under-recognized.[23] Risk factors for stent fracture are stent length (5cm vs. 3cm) and stents placed in bypass grafts or the RCA location.[24][25][26] The FDA requires manufacturers to demonstrate 10-year life duration for stents through stress testing.[27][26] Stent fracture without restenosis treated only conservatively shows good outcomes.[25] Malapposition is the problem of incomplete alignment of the stent struts to the vessel surface. It occurs in 2 to 5% of the cases and is a cause of late stent thrombosis.[28][29]

The Scaffold

Stents can be classified according to characteristics like metallic basis and design.[30] Different metallic compounds can serve as stent scaffold which gets further characterized by unique physical properties such as radial strength, ferromagneticity, material fatigue, and radiolucency.[31] Radial strength measures the resistance to external compression. Radiolucency characterizes the visibility of the stent during angiography.

Stainless steel or tantalum was used in early BMS.[31] Tantalum has a good visibility during angiography and lack of ferromagnetism. Following implantation, tantalum undergoes oxidation that makes it stable and resistant to degradation which results in better biocompatibility.[32][33] Tantalum showed similar thrombogenicity but better radio-opacity compared to stainless steel and is appreciated for its mechanical strength.[34][32] By comparison, 316L stainless steel with low carbon content is a mixture of iron, chromium, and nickel.

Later on, nitinol, a combination of nickel and titanium, was introduced into clinical practice as it exhibited elastic properties and thermal shape memory.[35] The thermal memory allows for the expansion of the nitinol stent once inside the body and smooth adaption to surrounding geometry.

Recent stent backbones are made from chromium combinations with either cobalt or platinum as they allow thin strut structure. Cobalt chromium stents are a choice in higher-risk lesions due to their high radial strength, low profile and increased elasticity.[36][37] Cobalt chromium scaffolds (MP35N, L605) compared with stainless steel scaffolds showed a lower risk for target vessel revascularization.[38][39][40] The platinum chromium stent possesses characteristics of good flexibility, deliverability, conformability, radial strength, and visibility.[41]

Due to improved scaffold materials having higher radial strength, strut thickness decreased from around 140 micrometers in thick-strut stents to around 70 micrometers in thin-strut stents. The ISAR-STEREO trial showed the superiority of thin-strut stents with lower restenosis.[42][43]

The problem of metal allergy plays no role in the pathogenesis of restenosis, but its influence on outcome shows contradictory evidence.[44][45][46] Some authors propose patch testing prior to implantation. [45] Nickel is especially known for its allergic potential and has been banned from textiles in many countries. Stainless steel, cobalt, chromium, and platinum chromium are the most commonly used alloys.[27]

Apart from the metallic basis, scaffold design has clinical implications as relates to vessel injury during implantation, and different designs result in different injury patterns with greater injury leading to increased tissue proliferation and platelet adhesion. [47] Stent designs include delta wing stent,[48] coil stent, tubular mesh stent, tubular slotted stent, and corrugated ring stent. The corrugated ring stent design showed less tissue proliferation than the tubular slotted stent design.[49]

Multicellular stents can subdivide into closed-cell and open-cell designs. Closed-cell stents show a more uniform drug delivery but are less flexible, whereas open-cell stents have a non-uniform drug distribution but improved alignment to the vessel.[50][51] In the PAST study, closed-cell stents triggered less platelet adhesion.[47] Uniform endothelial coverage can be achieved with flexible stent design and reduces stent thrombosis.[52]

Bioabsorbable stents are non-permanent implants avoiding caging vessels and obstructing side branches and allow vasomotion.[48] They are made either from metallic or polymeric scaffolds. The basis of metallic biodegradable scaffolds is magnesium or iron. The pharmacological properties of these two compounds are well known.[53] A growing number of synthetic or natural polymers are available to create biodegradable scaffolds such as poly-(L-lactic acid) (PLLA) or amino acids such as tyrosine.[54][53] 

The Therapeutic Agent

Antiproliferative drugs stop vascular smooth muscle cell proliferation and thus, neointimal hyperplasia. Sirolimus and paclitaxel were used in first generation DES. Paclitaxel inhibits microtubule disassembly and thus interferes with the cell cycle, leading to cell cycle arrest in G0-G1 and G2-M phases.[55] Sirolimus binds to the FKBP12 and subsequently inhibits mTOR and PI3 pathway, arresting the cell cycle in the G1 phase.[56][57] In clinical studies (i.e., REALITY, SIRTAX, ISAR-DESIRE) sirolimus-eluting stents (SES) outperformed paclitaxel-eluting stents (PES) regarding restenosis rate.[58] Although both inhibit the cell cycle differences in dosage, release kinetics, immunosuppressive properties, and distribution among the vessel wall layers might contribute to the clinical inferiority of paclitaxel-eluting stents.[59][60] Additionally, shear stress has been shown to vary between these two stent types.[61]

Therefore, later generation stent therapeutic agents evolved from sirolimus. Changes to the sirolimus structure created compounds such as everolimus (SDZ RAD) which showed anti-arteriosclerotic features and prevented graft rejection.[62][63] Ridaforolimus was non-inferior to zotarolimus in the BIONICS trial.[64] Further rapamycin agents include biolimus and novolimus. Concerns about tissue factor increase as a side effect of rapamycin agents have related to stent thrombosis.[65]

Other investigated therapeutic agents are[66][67][68][69][70][71][72][73][74]

• Dual drug-eluting stents DDES, which carry two therapeutic agents to combine their mechanism of action, for example antiproliferative and additional anti-thrombotic effects. The modification of the coating allows time controlled release of the drugs targeting the different timing of bio-response to stent implantation.
• Phytoncide (PTC), which shows the same anti-inflammatory and antiproliferative effects in vitro as sirolimus and thus might serve as an alternative to sirolimus
• Glucocorticoids, which are useful to suppress inflammatory changes that lead to restenosis
• Gene eluting stents, which deliver plasmid DNA to express appreciated proteins inside cells. The expression of VEGF, for example, supports healthy endothelialization of the stent luminal surface
• Galangin, which up-regulates p27KIP1 that arrests cell cycle in the G0-G1 phase, inhibits proliferation of vascular smooth muscle cells
• Tacrolimus reduces restenosis via the calcineurin/NFAT/IL-2 pathway but is inferior in head-to-head comparison to other stents eluting rapamycin agents
• Stem-cell carrying stents to support healthy reendothelialization
• Radioactive stents
• Actinomycin
• Probucol
• 7-Hexanoyltaxol

The Coating

Stent coating serves different purposes. It facilitates drug stent adhesion, drug-release, biocompatibility, and modulates thrombogenicity. Polymers are repeating sequences of chemical connections. Drug delivery via polymers follows the principle of diffusion, dissolution or ion exchange.[75] The group of coating materials further subdivide as[31]:

• Organic or inorganic
• Bio-erodable or permanent
• Uniform or nonuniform drug delivery
• Luminal or abluminal coating
• Active or passive

First generation DES used synthetic polymers such as poly(ethylene-co-vinyl acetate) (PEVA) and poly(n-butyl methacrylate) (PBMA) or tri-block copolymer poly(styrene-b-isobtylene-b-styrene) (SIBS). PEVA and PBMA coat SES whereas SIBS coats PES. SIBS allows early burst release of the drug which is ideal for controlling the early vascular injury following stent implantation. Observations of late stent thrombosis with PEVA and PBMA triggered investigations to find the underlying causes. Hypersensitivity arose as a factor as tissue surrounding stents became infiltrated by inflammatory cells.[76] To reduce inflammation following stenting, more biocompatible polymers such as phosphorylcholine (PC) and copolymer poly(vinylidene flouride-co-hexaflouropropylene) (PVDF-HFP) coat second-generation drug-eluting stents. Phosphorylcholine (PC) is a natural constituent of the phospholipid bilayer cell membrane. Since PC reduces platelet adhesion, it is useful to reduce thrombosis. Large vessel grafts are made from poly(bis(trifluoroethoxyl)phosphazene (PTFE).

Other coating options include[77][78]:

• Gold-coated stents, which exhibited an increased risk for restenosis
• Heparin
• Endothelial cells or its progenitors
• CD34 antibodies
• Natural glycocalyx

To further reduce the inflammatory response, stent coatings employed bioabsorbable polymers such as poly-lactic acid (PLA), poly(lactide-co-glycolide) (PLGA) and polycaprolactone (PCL).[79][80]  Experience with PLA derives from the application in clinical practice as sutures or screws and grafts and meshes.[81][82] A study comparing PLGA and PCL associated their different degradation to acidification of the surrounding tissue and cell proliferation and migration.[83] Poly(lactic acid-co-glycolic acid) (PLGA) degrades by hydrolysis in 6 months. It exhibits low immunogenicity, good biocompatibility, and mechanical characteristics.[84] Different trials (LEADERS, COMPARE II, NEXT, and CENTURY II) showed promising results for biodegradable polymer thin-strut drug-eluting stents compared with thin-strut permanent polymer drug-eluting stents.[85][86] 

The problems associated with coating polymers such as inflammation or hypersensitivity can be avoided altogether by using drug reservoirs stents or nanoparticles.

The plentitude of different drug-eluting stents requires knowledge about their characteristics to align the advantages and disadvantages of the stent with the requirement of each unique lesion. Stent performance depends strongly on lesion characteristics.[87] Arterial response to stent implantation is complex, and thus preclinical (i.e., animal) studies are key to predict performance in humans.[88] There is not one ideal stent, but different stents are better suited for different situations. For example, small diameter stents have been designed for narrow vessel treatment (<2.5 to 3.0mm).[89][90] Since real-world experience showed frequent off-label usage, there has been a broadening of the indications for DES placement.[91] 

New-generation DES can be used in complex lesions such as with diabetic patients,[92] chronic kidney disease,[93] acute myocardial infarction (as was investigated in the HORIZONS-AMI trial),[94] ostial lesions or bifurcations,[95] bypass grafts, as an alternative to coronary artery bypass grafting in left main stem disease,[96] small vessels or long lesions, restenosis, and chronic occlusions.

Bare metal stents are not outdated, but drug-eluting stents have largely replaced them.[39] Some authors argue that practice adaption and application of DES has been too fast.[14] BMS are still used in about 20% of PCI and indicated in patients that have limited DAPT options (awaiting surgery, compliance issues, bleeding risk, preexisting atrial fibrillation). BMS may also be preferable in large vessel lesions, acute myocardial infarction, and regarding costs.[97] In the US, a BMS costs round about 1,000 USD and a DES around 3,000 USD. Comparing costs requires considering long-term outcomes that favor DES because of less reintervention and thus improved cost-benefit-analysis.[98][99][100]

The discussion about the appropriateness of real-world stenting practice is ongoing. For example, the NORSTENT study emphasized the advantages of BMS compared to DES. Designing studies and finding evidence is difficult, which can be illustrated by the fact that the progression of restenosis differs, which makes early angiographic measures unrelated to late stenotic changes. Measuring restenosis risk as a surrogate is common among studies, but it may not correlate with clinical endpoints such as death and myocardial infarction.[101] 

Evidence advocates for everolimus-eluting stents, which exhibit the most favorable outcomes and the future might belong to biodegradable polymers and bioabsorbable scaffolds.[102][103] Until now evidence did not show the superiority of bioresorbable scaffold stents over second-generation drug-eluting stents.[48] Lee and Torre Hernandez describe in their article the ideal stent of the future as showing good deliverability, having thin struts, being highly visible, carrying a rapamycin agent that is eluted for two to three months and being coated by a thin layer polymer allowing for short duration DAPT.[104]