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British Medical Bulletin 59:227-248 (2001)
© 2001 Oxford University Press
Stent development and local drug delivery
Department of Cardiology, Thoraxcentre, Erasmus Medical Centre Rotterdam, The Netherlands
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Abstract |
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Abstract
IntroductionStent implantation has become the new standard angioplasty procedure. In-stent re-stenosis remains the major limitation of coronary stenting. Re-stenosis is related to patient-, lesion- and procedure-specific factors.
Patient-specific factors can not be influenced to any extent. Procedure-specific factors are affected by implantation technique and stent characteristics. Design and material influence vascular injury and humoral and cellular response. Radiation has been shown to have inhibitory effects on smooth muscle cell growth and neo-intima formation, but in clinical trials the outcome has been hampered by re-stenosis at the edges of the radioactive stent (‘candy wrapper’).
New approaches target pharmacological modulation of local vascular biology by local administration of drugs. This allows for drug application at the precise site and time of vessel injury. Systemic release is minimal and this may reduce the risk of toxicity. The drug and the delivery vehicle must fulfil pharmacological, pharmacokinetic and mechanical requirements and the application of eluting degradable matrices seems to be a possible solution. Numerous pharmacological agents with antiproliferative properties are currently under clinical investigation, e.g. actinomycin D, rapamycin or paclitaxel. Another approach is for stents to be made of biodegradable materials as an alternative to metallic stents. Their potential long-term complications, such as in-stent re-stenosis and the inaccessibility of the lesion site for surgical revascularization, needs to be assessed.
Current investigational devices and the line of (pre)clinical investigation are discussed in detail. Currently, there is little experimental, and only preliminary clinical, understanding of the acute and long-term effects of drug-eluting or biodegradable stents in coronary arteries. The clinical benefit of these approaches still has to be proven.
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Introduction |
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Over the last decade, coronary stents have revolutionized the field of interventional cardiology. Stent implantation has become the new standard angioplasty procedure1
–3. This popularity is mainly for two reasons: (i) the unique capability to master a major complication of balloon angioplasty – (sub)acute vessel closure; and (ii) a superior long-term outcome in comparison to balloon angioplasty4–8. The high reliability of the acute angioplasty result after stenting allowed for a dramatic expansion in the indication for catheter-based intervention (ostial lesions9, bifurcation lesions10,11, left main lesions12,13, multiple lesions14). In the early phase, device technology was directed toward improving the accessibility of lesions, e.g. in tortuous, small and/or calcified vessels. The advances included premounted stents, reduced crossing profile and improved flexibility. Modifications in stent design such as ‘rotating’ and ‘locking’ mechanisms afforded them high flexibility when unexpanded and remarkable radial strength when expanded. Further development included specifically configured stents for distinct indications, such as bifurcation lesions, lesions with side branch take-off, aneurysms and vessel rupture.
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The problem: re-stenosis |
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As most acute aspects of stenting are resolved, the long-term outcome becomes more and more the focus of attention. In-stent re-stenosis remains for several reasons the major limitation of coronary stenting. The absolute number of in-stent re-stenotic lesions is increasing in parallel with the steadily increasing number of stenting procedures and with the complexity of culprit lesions. The treatment of in-stent re-stenosis is, despite progresses in radiation therapy, technically challenging and costly. In subsets of lesions (such as small vessel size and diffuse disease), an anticipated high risk for re-stenosis may prevent the use of stents.
Re-stenosis is considered as a local vascular manifestation of the general biological response to injury. Catheter-induced injury consists of denudation of the intima and stretching of the media. The wound-healing reaction consists in an inflammatory phase, a granulation phase and a remodelling phase. The inflammatory phase is characterized by platelets and growth factor activation, the granulation phase by smooth muscle cell and fibroblast migration and proliferation into the injured area, the remodelling phase by proteoglycan and collagen synthesis, which replaces early fibronectin as major component of extracellular matrix15. Current concepts describe three mechanisms of the re-stenotic process: early elastic recoil, late vessel remodelling and neo-intimal growth16,17. Coronary stents provide mechanical scaffolding that virtually eliminates recoil and remodelling18. However, neo-intimal growth continues to be a major problem.
Neo-intimal proliferation occurs principally, but not exclusively, at the site of the primary lesion within the first 6 months after stent implantation. Neo-intima is basically an accumulation of smooth muscle cells within a proteoglycan matrix that narrows the previously enlarged lumen. Neo-intima formation is triggered by a cascade of cellular and molecular events including platelet activation, leukocyte infiltration, smooth muscle cell expansion, extracellular matrix elaboration and re-endothelialization.
In the clinical setting, neo-intimal formation and in-stent re-stenosis are related to patient-specific factors such as genetic predisposition or diabetes mellitus19, to lesion-specific factors such as vessel calibre20, lesion length or plaque burden 21 and to procedure-specific factors such as extent of vessel damage, residual dissections22, number of stents, minimal stent diameter or minimal stent area23.
Patient-specific factors can not be influenced to any extent. Attempts to modulate lesion-specific factors by pharmacological therapy have not been not successful so far24,25. Procedure-specific factors may be favourably affected by the stent implantation technique and stent characteristics26.
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The solution: the ideal stent? |
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From the interventionist's technical point of view, the ideal stent should have at least the following features:
- • Reach any location within the coronary artery system
- • Have a low profile, high trackability and flexibility
- • Allow for precise placement with good visibility
- • Guarantee predictable expansion with minimal foreshortening
- • Provide equally distributed radial strength, be rigid and have many, thick struts
- • Prevent plaque protrusion having many struts with small spaces between struts
- • Allow for side-branch access having few struts with large spaces between struts
- • Reduce acute thrombotic and inflammatory vessel response, i.e. have few, thin struts
- • Allow wound healing but inhibit exaggerated neo-intimal growth. It should be thromboresistant and have high hydrodynamic compatibility
- • Guarantee vessel accessibility for further intervention (PCR, bypass grafting) and thus possibly disappear after 6 months.
Unfortunately, all these requirements on stent characteristics are mutually incompatible.
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The ideal stent: small steps towards a big goal |
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The focus of stent development in the first decade of stenting was on optimization of stent characteristics. Systematic investigations were performed to gain insights into mechanisms of stent action and vessel biology. This resulted in a huge variety of stents, differing in design, material, surface, radioactivity and coatings27
. Design Stent design varies in the geometry (number of intersections and interstrut area), in the strut configuration and the metal-to-artery ratio. These are the major determinants of stent profile, flexibility, radial strength and (elastic) expansion characteristics.
Presently, more than 55 stents are available (12 with FDA-approval) which are manufactured by more than 30 companies. They use two different expansion principles (balloon-expansion and self-expansion) and can be categorized in five basic design types: tubular, ring, multi-design, coil, and mesh.
The physical properties affect vascular injury and response. The different stent designs show a considerable range of the elastic moduli28 which affect stent expansion properties as late as 8 weeks after deployment29. Reduction of strut–strut intersections can reduce the vascular injury score, thrombosis and neo-intimal hyperplasia30. Non-uniform stent expansion increases vascular injury and in-stent re-stenosis31.
Material and surface (texture)
Stent materials and surface are relevant to in-stent re-stenosis. Surface characteristics can be modulated by (electromechanical) polishing, ion implantation or coating. Suitable stent materials are metals, metal-alloys or polymers. Chemistry, charge, and texture32 modulate humoral and cellular vessel response (plasma proteins, inflammatory and proliferative mediators, platelet and leukocyte activation). Copper is more frequently associated with subacute thrombosis than steel33.
Radioactivity
Another possible way to modify the physical properties is the introduction of radioactivity34,35. Radiation has proven inhibitory effects on smooth muscle cell growth and neo-intima formation36. However, clinical trials utilizing radioactive stents have been disappointing. Despite effective prevention of neo-intimal growth with the stent, clinical and angiographic outcomes have been hampered by re-stenosis at the edges of the radioactive stent; termed the ‘candy wrapper’ effect (Fig. 1)37. This unfavourable phenomenon occurred irrespective of the stent design (cold end, hot end) or the dose rate (high activity versus low activity; Table 1).
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Coating Coating categories are various (Table 2). Stent coatings can dramatically reduce protein deposition and platelet adhesion in experimental settings (Fig. 2)38
. However, polymers have shown conflicting results in the experimental setting39–41 with some provoking a severe tissue response42. A number of other coatings like inert polymer30, phosphorylcholine43,44 or heparin45–47, demonstrated a reduction in (sub)acute stent thrombosis rate and possible effects on neo-intimal hyperplasia48.
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In clinical practice, however, the acute beneficial effect on stent thrombosis is of minor relevance as already modern uncoated stents show a very low (sub)acute thrombosis rate. Furthermore, the acute beneficial effect did not result in a substantial decrease in in-stent re-stenosis49
. In response to this, the interest in coatings has shifted towards considering coatings as vehicles for local drug delivery. So far, phosphorylcholine is the only clinically-available, polymer coated stent (Fig. 3)50 .
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A big step: local drug delivery? |
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New approaches target not only the stent characteristics but also the pharmacological modulation of the local vascular biology. A proposed explanation for the repeated failure of clinical drug studies has been that agents given systemically cannot reach sufficient levels in injured arteries to impact significantly on the re-stenotic process. Local administration of drugs offers advantages. The active drug is applied to the vessel at the precise site and at the time of vessel injury. Local drug delivery might be able to achieve higher tissue concentrations of the drug. No additional materials or procedures are required. Systemic release is minimal and may reduce the risk of remote or systemic toxicity.
Candidate delivery vehicles The delivery vehicle must fulfil pharmacological, pharmacokinetic and mechanical requirements. The release of the drug into the vessel must take place in a manner that is consistent with the drug's mode of action. Drug-release must be in predictable and controllable concentrations and within a known time span. The delivery vehicle must be suitable for sterilisation. It must follow the stent changes of configuration during stent expansion and resist mechanical injury caused by the implantation balloon. An example of possible deleterious effects of stent expansion and sterilisation on the coating is given in Figure 4. Currently, these problems are controlled, guaranteeing intact coating during clinical application (Fig. 5). The application of eluting, degradable matrices seems to be a possible solution. An overview of delivery vehicles for drug eluting systems is given in Table 3.
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Candidate drugs The drug should be able to inhibit the multiple components of the complex re-stenosis process. Uncontrolled neo-intima tissue accumulations shows some parallels to tumour growths; thus, the use of anti-tumour strategies seems to be a logical choice. Numerous pharmacological agents with antiproliferative properties have been tested for their potential to inhibit re-stenosis with mostly disappointing results51
. Antimitotic compounds (like methotrexate and colchicine) have failed to inhibit smooth muscle cell proliferation and intimal thickening52,53. In contrast, other agents such as angiopeptin54, GP IIb/IIIa inhibitors or steroids5556 have shown a promising inhibitory effect on neo-intimal proliferation. Potential candidates for local drug delivery are given in Table 4. The following drugs are now being tested in randomized clinical trials.
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Actinomycin D (Cosmegen®)
Actinomycin D has been marketed world-wide since the 1960s. It is an antibiotic used for its antiproliferative properties in the treatment of various malignant neoplasmas (e.g. Wilms tumour, sarcomas, carcinoma of testis and uterus). It inhibits the proliferation of cells. Actinomycin D (C62H86N12O16) forms, via deoxyguanosine residues, a stable complex with double-stranded DNA and inhibits DNA-primed RNA synthesis.
Rapamycin (Sirolimus; Rapamune®)
Rapamune® is an FDA approved drug for the prophylaxis of renal transplant rejection in use since 1999. It is a naturally occurring macrocyclic lactone (Fig. 6) which is highly effective in preventing the onset and severity of disease in several animal models of autoimmune disease, such as insulin-dependent diabetes mellitus, systemic lupus erythematosus and arthritis.
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The class of macrocyclic immunosuppressive agents (rapamycin, cyclosporin A, tacrolimus) bind to specific cytosolic proteins called immunophilins to gain their immunosuppressive activity. Rapamycin blocks G1 to S cell cycle progression (Fig. 7) by interacting with a specific target protein (mTOR, mammalian target of rapamycin) and inhibiting its activation. The inhibition of mTOR suppresses cytokine-driven (IL-2, IL-4, IL-7 and IL-15) T-cell proliferation.
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mTOR is a key regulatory kinase and its inhibition has several important effects including: (i) inhibition of translation of a family of mRNAs that code for proteins essential for cell cycle progression; (ii) inhibition of IL-2-induced transcription of proliferating cell nuclear antigen (PCNA) that is essential for DNA replication; (ii) blocking CD28-mediated sustained up-regulation of IL-2 transcription in T cells; and (iv) inhibition of the kinase activity of the cdk4/cyclin D and cdk2/cyclin E complexes, essential for cell cycle progression. On overview of rapamycin effects within the cell cycle is given in Figure 8.
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The mechanism of action is distinct from other immunosuppressive drugs that act solely by inhibiting DNA synthesis, such as mycophenolate mofetil (CellCept) and azathioprine (Imuran). Rapamycin is synergistic with cyclosporin A and has much lower toxicity than other immunosuppressive agents.
In in vitro and in vivo studies, rapamycin prevents proliferation of T cells but also proliferation57,58 and migration59 of smooth muscle cells. Furthermore, rapamycin has been shown to diminish smooth muscle cell hyperproliferation in several animal models of arteriopathy60–62.
Paclitaxel (Taxol®)
Paclitaxel was originally isolated from the bark of the Pacific Yew. It is an antineoplastic agent that is currently used to treat several types of cancer, most commonly breast and ovarian cancer.
It is a diterpenoid with a characteristic taxane-skeleton of 20 carbon atoms and has a molecular weight of 853.9 Da. Paclitaxel exerts its pharmacological effects through formation of numerous decentralized and unorganized microtubules. This enhances the assembly of extraordinarily stable microtubules, interrupting proliferation, migration and signal transduction63,64 Unlike other antiproliferative agents of the colchicine type, which inhibit microtubuli assembly, paclitaxel shifts the microtubule equilibrium towards microtubule assembly. It is highly lipophylic, which promotes a rapid cellular uptake, and has a long-lasting effect in the cell due to the structural alteration of the cytoskeleton.
In vitro and in vivo studies have shown that paclitaxel may prevent or attenuate re-stenosis. Paclitaxel inhibits proliferation and migration of cultured smooth muscle cells in a dose-dependent manner65. In a rat balloon injury model, intraperitoneal administration of paclitaxel reduced neo-intimal area. In a rabbit atherosclerotic model where plaque burden was increased by electrical injury, local administration of paclitaxel reduced neo-intimal thickness66,67.
Investigational devices
Actinomycin D: Multi-Link TetraTM-D stent (Guidant, Santa Clara, CA, USA)
The stent is fabricated from medical 316L stainless steel tubing and is composed of a series of cylindrically oriented rings aligned along a common longitudinal axis. Each ring consists of 3 connecting bars and 6 expanding elements. The stent is premounted on a delivery catheter.
The antiproliferative drug is actinomycin D. The finished Multi-Link Tetra stent is coated with a polymer matrix (semicrystalline ethylene-vinyl alcohol co-polymer: EVAL) containing a maximal dose of 150 µg actinomycin D. This is equivalent to 20–200 times less than the recommended total human adult dose of 500 µg/day given intravenously for 5 days.
The delivery catheter is a rapid exchange design (0.014 inch guidewire). It is equipped with two radiopaque markers located underneath the balloon to mark the ends of the stent and has a ‘stepped’ balloon design to optimize stent and balloon shoulder configuration.
NIRxTM – paclitaxel-coated conformer coronary stent (Boston Scientific, USA)
The stent is fabricated from medical 316LS stainless steel. The geometry is a continuous, uniform, multicellular design with adaptive cells capable of differential lengthening. This enables the stent to be flexible in the unexpanded configuration. Stent length is 15 mm. The stent is premounted on a delivery catheter (Fig. 9).
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The antiproliferative drug is paclitaxel. Paclitaxel is incorporated into a fast-release triblock co-polymer carrier system on the stent. There are two drug concentration. The ‘low dose’ concentration is 1.0 µg/mm2(loaded drug/stent surface area; total dose 85 µg per stent) and gives sustained release over
28 days. The ‘moderate dose’ is 2.0 µg/mm2 (loaded drug/stent surface area).and provides a rapid release in the first 24 h, followed by a slower release over the following 28 days. The delivery catheter is a monorail design (0.014 inch guidewire/7F guiding catheter). It is equipped with two radiopaque markers located underneath the balloon to mark the ends of the stent. The delivery balloon will be 3.0 mm and 3.5 mm in diameter.
Rapamycin-coated BXTM VELOCITY stent (Cordis, Warren, USA)
The stent is fabricated from medical 316LS stainless steel. It is available in 2 configurations: a 6-cell configuration (expanded diameter 2.5–3.25 mm) and a 7-cell design (expanded diameter 3.5–3.75 mm). Stent length is 18 mm. The stent is premounted on a delivery balloon with diameters of 2.5 mm, 3.0 mm and 3.5 mm (Fig. 10).
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The antiproliferative drug is rapamycin. The stent contains 140 µg/cm2 which gives a total rapamycin content of 153 µg on the 6-cell stent and 180 µg on the 7-cell stent. The coating formulation consists of 30% rapamycin by weight in a 50:50 mixture of the polymers polyethylenevinylacetate (PEVA) and polybutylmethacrylate (PBMA; Fig. 11).
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The delivery catheter utilizes a rapid exchange design (0.014 inch guidewire/7F guiding catheter). It is equipped with two radiopaque markers located underneath the balloon to mark the ends of the stent. The delivery balloon will be 3.0 mm and 3.5 mm in diameter.
Clinical studies
Actinomycin D
There is no published research to date documenting the use of actinomycin D for treatment of coronary artery disease and/or re-stenosis. A phase 1, randomized clinical trial ACTION (ACTinomycin eluting stent Improves Outcomes by reducing Neointimal Hyperplasia) started in June 2001 to evaluate the safety and performance of the Multi-Link TetraTM-D stent system: 360 patients will be randomized to receive an actinomycin D coated stent or a non-coated stent for treatment of de novo lesions in native coronary arteries with a vessel calibre of 3.0–4.0 mm. Six month angiographic follow-up is expected to be completed in February 2002, 12-month clinical follow-up up is expected to be completed in August 2002.
Rapamycin (Sirolimus)
A first clinical application of the rapamycin-coated stents was performed in Sao Paulo and Rotterdam. Thirty patients with angina pectoris were electively treated with 2 different formulations of rapamycin-coated BXTM VELOCITY stents (Cordis) (slow release [SR], n = 15, and fast release [FR], n = 15). All stents (18 mm) were successfully delivered, (3.0–3.5 mm vessel calibre) and patients were discharged without clinical complications. At 4 months angiographic and IVUS follow-up, there was minimal neo-intimal hyperplasia in both groups (11.0 ± 3.0% in the SR group and 10.4 ± 3.0% in the FR group, P = NS) by ultrasound and quantitative coronary angiography (in-stent late loss, 0.09 ± 0.3 mm [SR] and –0.02 ± 0.3 mm [FR]. No in-stent or edge re-stenosis was observed. No major clinical events (stent thrombosis, repeat revascularization, myocardial infarction, death) had occurred by 8 months68. At 1 year follow-up, IVUS volumetric analysis and angiography indicated minimal amounts of neo-intimal hyperplasia that were scarcely different from the 4 month data in both groups, with some patients showing no evidence of hyperplasia whatsoever. There were no MACE and no re-stenosis in either of the groups. One late acute MI occurred in the fast-release group at 14 months69.
The randomized RAVEL study with the rapamycin-coated BXTM VELOCITY balloon-expandable stent in the treatment of patients with de novo lesions in native coronary arteries is a multicenter, prospective, randomized double-blind clinical trial comparing bare metal and the drug-coated stents. A total of 220 patients were randomized for treatment with either a single rapamycin-coated (140 µg.cm–2) or a bare metal BXTM VELOCITY stent. At 6-month follow-up, the restenosis rate of the treated group was zero, the loss in minimal lumen diameter was zero, there was no target lesion reintervention and the event-free survival was 96.5%74.
The SIRIUS study is a multicentre, prospective, randomized double-blind trial that is being conducted in 55 centres in the USA. Eleven hundred patients with focal de novo native coronary arterial lesions (2.5 to 3.5 mm diameter, 15 to 30 mm long) will be randomized for treatment with either rapamycin-coated (109 mg.cm–2) or bare metal BXTM VELOCITY balloon expandable stents. The primary endpoints of the SIRIUS trial are target vessel failure (death, myocardial infarction, target lesion revascularization) at 9 months. In addition, secondary endpoints are core laboratory analysis of angiographic and intravascular ultrsound data to determine treatment effects on neointimal hyperplasia and in-stent restenosis. Clinical follow-up will continue for 3 years in order to assess late events. In addition to the pivotal RAVEL and SIRIUS trials, feasibility studies are ongoing to assess efficacy of rapamycin-coated stents in more complex lesion subsets, such as in-stent restenosis.
Paclitaxel
There are several ongoing clinical trials of paclitaxel-coated stents. In the TAXUS I trial 61 patients were randomized to receive a paclitaxel-coated (1.0 µg/mm2) or a bare NIR stent. At 6-month follow-up, no restenosis was seen in the paclitaxel-coated stent group, while the restenosis rate in the bare stent group was 11%. The late lumen loss of 0.35±0.47 mm was significantly lower in the paclitaxel-coated stent group (0.71±0.88 mm). The Asian ASPECT trial showed a clear dose response. Patients (n= 177) were randomized to receive a high dose (3.1 µg.mm–2) paclitaxel-coated, a low dose (1.3 µg.mm–2) paclitaxel-coated, or a bare stent. The restenosis rate at 6 months was 4%, 12%, and 27%, respectively75. The ongoing ELUTES trial randomises 180 patients into 5 groups – 4 different dose levels and a bare stent control group.
In other clinical trials, the taxol derivate QP2 was used. The QP2 pilot study included 32 patients with de novo or restenosis lesions who underwent QuaDS-QP2 stent implantation. The stainless steel, slotted tube stent was 13 or 17 mm in length and coated with multiple ploymer sleeves that slowly release QP2 (up to 4000 µg). A 2-year follow-up of 25 patients recorded that they were all asymptomatic76. Another group reported on the 8 month IVUS follow-up. IVUS revealed only moderate neointima formation with a neointima burden of 13.6±14.9%77. Another multicenter, trial randomized 266 patients to receive a QuaDS-QP2 stent (4000 mg with an elution over 180d) or a bare stent. Follow-up angiography at 6 months showed a significant reduction in restenosis in the QuaDS-QP2 stent group (6.9% v. 36%). However, this trial has been stopped by the safety committee because of an excessive adverse event rate in the QuaDS-QP2 stent group of 10.2% – periprocedural myocardial infarction and subacute stent thrombosis78.
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Another step: biodegradable stents |
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Coronary stents exert their beneficial clinical effect within a relatively narrow time frame. Stent scaffolding is needed from the acute procedural phase (in case of threatened or actual vessel closure) until the first 6 months after the procedure to overcome late negative vessel remodelling. In the long-term perspective, metallic stents have potential complications such as in-stent re-stenosis and the inaccessibility of the lesion site for surgical revascularization. Thus stents made of biodegradable materials may be an ideal alternative.
Experimental data The first requirement for polymeric materials in intracoronary stents is biocompatibility. In animal studies, the biocompatibility of polymer stents has been controversial suggesting that tissue incompatibility may be a major obstacle.
Marked inflammatory response after the implantation of 5 different polymer-loaded stents (polyglycolic acid/polylactic acid, polycaprolactone, polyhydroxybutyrate valerate, polyorthoester, and polyethyleneoxide/ polybutylene terephthalate) has been reported in a porcine coronary model42. Thrombotic occlusion of polymeric stents was seen in other experiments70.
In contrast, Zidar et al reported only a minimal inflammatory reaction and minimal neo-intimal hyperplasia with the use of poly-l-lactic acid (PLLA) stents in canine femoral arteries. In vitro data revealed a reduced platelet adherence and thrombogenicity of the PLLA stent as compared with slotted-tube stainless steel metallic stents70.
Investigational device: poly-l-lactic monopolymer Igaki-Tamai biodegradable stent (Igaki Medical Planning Co, Ltd) The Igaki-Tamai stent is a coil stent made of a poly-l-lactic (PLLA) monofilament (molecular mass, 183 kDa). PLLA has been used for orthopaedic applications in humans and has generally been found to be biocompatible. The stent is self-expanding with a zigzag helical design. The stent length is 12 mm. The thickness of the stent strut is 0.17 mm. In its expanded state, the stent covers 24% of the vessel area. The stent has a radiopaque gold marker at both ends of the prosthesis. Stents are mounted on standard angioplasty balloon catheters that are the same size as the stent with diameters of 3.0, 3.5, and 4.0 mm (Figure 12). It takes the stent 18–24 months to biodegrade fully.
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Deployment of the stent is currently done with a balloon-expandable covered sheath system through an 8 French guiding catheter. The stent delivery balloon inflation is performed with a heated dye at 80°C using a 30 s inflation at 6–14 atm. This temperature ensures adequate stent expansion within 30 s and may minimize vessel injury caused by a heated balloon. The stent continues to expand gradually to its original size after deployment in vivo.
Clinical data
The first clinical data are available in 15 patients who underwent elective coronary PLLA Igaki-Tamai stent implantation: 25 stents were successfully implanted in 19 lesions. Angiographic success was achieved in all procedures. No stent thrombosis and no major cardiac event occurred within 30 days. Angiographically, both the re-stenosis rate and target lesion revascularization rate were 10.5% at 6 months. Intravascular ultrasound revealed no significant stent recoil at 1 day and stent expansion at follow-up. No major cardiac event, except for repeat angioplasty, developed within 6 months71.
Long-term data (12 months) of 63 lesions in 50 patients have been previously presented. Angiographic analysis showed a good procedural result with was a 12 ± 8% final diameter stenosis post-stent implantation. At 6 months, diameter stenosis was 38 ± 23% and decreased slightly to 33 ± 23% at 12 months. Subacute thrombosis occurred in one patient at day 5, but no other MI, urgent CABG, or death occurred over the 12 months72.
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Limitations |
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Although the principle of stent implantation is well established and although (most of) the applied drugs and polymers have been used in clinical practice for many years, there is little experimental and only preliminary clinical knowledge of the acute and long-term effects of drug-eluting or biodegradable stents in coronary arteries. Thus, a number of concerns and open questions have to be investigated in the future.
The concerns include drug toxicity as well as acute and late vascular effects. A number of toxic effects are known for all drugs applied in cancer therapy such haematological toxicity (neutropenia), neurotoxicity (peripheral neuropathy), hypersensitivity reactions, or cardiac disturbances. However, these side-effects are described in patients undergoing high-dose chemotherapy for a malignant disease with plasma levels 100–1000 times higher (and over longer time periods) than plasma levels that result from a local delivery. Another concern is possible delayed wound healing and endothelialization. This would increase thrombogenicity and the danger of (late) stent thrombosis. Further potential side-effects could be late positive remodelling and aneurysmal formation. Thus the most suitable antithrombotic regimen following drug-eluting and/or biodegradable stent implantation still has to be evaluated.
A series of open questions exists on the mechanism of action and also the design of local drug delivery systems and drug-eluting stents. Little is known of specific pharmacokinetic issues. There is a paucity of data on the most appropriate tissue concentration and the rate and duration of drug-release over time. The tissue concentration is dependent on close mechanical contact of the stent to the vascular tissue and on physiological transport forces into the tissue. Hydrophobic drugs, like paclitaxel, have greater variability in terms of drug delivery, while hydrophilic drugs, like heparin, have less variability and achieve higher local concentrations. Local concentrations and concentration gradients, however, are crucial parameters for biological effects. The relationship between vascular effects and physicochemical properties of the drug-loaded stent is poorly understood. Drug distribution within the vessel wall seems to be significantly affected by the stent expansion pattern (uniform versus non-uniform)73.
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Conclusions and future perspectives |
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Drug-eluting and biodegradable stents represent one of the fastest growing fields in interventional cardiology today. However, many unanswered questions still have to be resolved before determining the potential of these techniques. Hopefully, after the completion of planned and on-going trials many of these issues will be answered. Furthermore, these new technologies will have to prove effective in the daily routine of treating patients presenting with long lesions, small vessels, chronic occlusion, bifurcation, multi-vessel and/or left main stem disease or acute myocardial infarction.
Stent development will investigate a variety of possibilities to resolve the re-stenosis problem. Possibilities range from the further exploitation of different classes of drugs which are potential candidates for the inhibition of re-stenosis to the combination of biodegradability with drug delivery, local gene therapy (e.g. local expression of proliferation regulatory genes; transfer of cytotoxic genes, VEGF) or external heating of stents (Fig. 13).
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Acknowledgements |
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ER was supported by a grant from the Deutsche Forschungsgemeinschaft.
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Footnotes |
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Correspondence to: Prof. P W Serruys, Thoraxcentre (Bd 406), Erasmus Medical Centre Rotterdam, Dr Molewaterplein 40, 3015 GD Rotterdam, The Netherlands
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