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Drug-Eluting Bioresorbable Stents


Dr Xiang Zhang and Phil Jackson of CERAM look at the material options, including different polymer combinations for next generation bioresorbable stents.  

A new era for coronary stent technology started at the beginning of 2011. Abbott obtained the CE-mark for ABSORB,1 the world’s first drug-eluting bioresorbable stent for the treatment of coronary artery disease. The device restores blood flow by opening a clogged vessel and provides support to the vessel until the device dissolves within approximately two years, leaving patients with a treated vessel free of a permanent metallic implant. The new stent is made of biodegradable polylactide (PLA). This polymer material was developed in 1960 and is important for many applications, including sutures, screws, pins/rods, drug delivery systems and now coronary stents.

The question is, "Given that metallic stents have been extensively developed and are on the market, why do we need biodegradable stents?” The answers are varied and include the following considerations.
  • There is no need for permanent coronary scaffolding beyond the first 6 to 12 months when the process of intimal hyperplasia and acute chronic recoil is completed.
  • Thirty percent of restenosis occurs when bare metal stents are used. This and other associated problems have not disappeared with drug-eluting stents (DES), despite dual antiplatelet agents being employed.2
  • Excessive use of metal stents may interfere with traditional re-interventional techniques such as bypass graft surgery.
  • Metal stents may pose problems with modern imaging techniques such as magnetic resonance imaging and multi-slice computerised tomography.3
Developing the biodegradable stent
There is no doubt that bioresorbable stents have clear advantages over bare metal and drug-eluting stents. However, we are just at the start of the journey because the technology is far from mature.

For stents, as for other medical devices, adverse incidents are often caused by inadequate materials being employed, whether these are bare metal, drug-eluting or bioresorbable stents. We have seen improvements in this area, but unfortunately not complete and satisfactory solutions yet. From a development point of view, it is vital to get the materials right first time before commencing design control, because of the high cost and the long time it takes to develop medical implants. The material issue is often not well addressed initially; rather more attention is paid to it during the later stage of clinical trials.

Before discussing stent and associated materials technology, it is important to appreciate the composition of vessels carrying blood and the mechanisms behind plaque formation.

Mechanics of plaque formation
Arteries and veins comprise three layers when the walls are viewed in cross-section. An outer layer, the adventitia, which encloses a middle layer of smooth muscle cells (SMC) called the media; and an inner layer, the intima, which contains endothelial cells. The endothelial cells become damaged with time. Platelets and white blood cells then collect at these cells. Eventually cells in the media layer burst through the intima and mix with intima cells in the artery interior. Fat from the blood then mixes in. The plaque formed can rupture with blood platelets collecting to form a clot that ultimately causes a heart attack.

Other stent materials technology
The first commercialised stents were bare metal stents, which appeared around 1994. Some of the metals employed were nitinol (nickel-titanium) alloys, which exhibit a combination of strong and flexible properties that are particularly suited to self-expanding stents. More commonly, stainless steel and more recently cobalt–chrome alloys are employed. One problem associated with metal stents is that they can trigger an inflammatory response and the growth of SMC, which causes restenosis; this can occur in as many as 30% of cases.4

To overcome the restenosis problem associated with bare metal stents, DES have been available since 2000. Early DES development work focused on characterising the eluting profile associated with cytostatic drugs mixed in a polymer coated on a metal stent.5 The cytostatic drugs have been reported to bring about a dramatic reduction in the occurrence of restenosis.

In general, polymer coatings provide an initial barrier over the metal and dissolve slowly to release drugs that retard restenosis. Polymeric materials as stent coatings have clear advantages, not least the ability for relatively low temperature processing that allows the introduction of organic drug molecules and even cells. Examples of the latter include introducing progenitor cells (that encourage endothelialisation around inserted stents) or growth factor (protein) compounds that retard SMC formation caused by disturbing the endothelial layer. However, longer-term issues with blood clotting can arise, possibly associated with SMC formation following dissolution of the polymer layer and re-exposure to metal.

Soluble glass and bioceramic
As alternatives to polymers, ceramic and glass materials have the potential to act as drug release vehicles for DES applications. CERAM has recently developed particulate inorganic drug release entities via water-soluble glasses and sol-gel materials. Both systems are versatile and can be formulated to deliver a range of release profiles versus time. In glass systems, the ratio of "network formers” (such as Si-O or P-O) to "network modifiers” dictates durability in an aqueous environment. By using created glasses that fully dissolve in water, options for adding temperature-sensitive drugs are introduced. Sol-gel technology offers a further route for low temperature preparation of drug/inorganic composites. By starting with alkoxide molecules, for example, Si(OR)4 for silica, glass-like structures can be built "bottom-up” via hydrolysis and condensation reactions.

The above reactions can occur around dissolved drug material so that a composite monolith forms on drying. Depending on the drug chemistry, the potential exists for silanol groups to react with a functional group in the drug. Final sol-gel structures can be tailored to deliver micro- or meso-porosity and great versatility in terms of drug release via diffusion from pores. It is possible for the aforementioned inorganic drug-carrying systems to be employed as coatings on metal stents with a better controlled release profile. Figure 1 illustrates that drug release rates can be controlled through formulation. The data relates to powdered materials crushed from glass/drug and sol-gel/drug monoliths. It demonstrates that drug release rate can be designed to be very slow or very fast (and everything in between) for the same drug.  For glass systems, dissolution according to glass chemistry is essential. For sol-gel systems, diffusion through different porous structures is critical, porosity can be controlled via method of catalysis and drying rates. 

Figure 1: Effect of formulation and process of sol-gel glass on drug release rate

Bioresorbable stent development To make an ideal drug-eluting bioresorbable stent (DEBS), the crucial step at the start is material selection. As a minimum, the following material characteristics should be considered:
  • micro mechanical properties and micro structure to meet the performance requirements
  • biocompatibility and bioactivity that help to cure the diseased coronary artery quickly
  • material bioresorbablity, that is, the time it takes to do the job and "disappear,” with the capability of being tailored to suit each patient’s individual needs
  • drug-eluting concentration and the time required for the curing process.
Micro mechanics and micro physics
It is essential to consider the physical structure and mechanical properties of a blood vessel. Figure 2 shows a blood vessel hoop stress as a function of blood pressure and internal diameter. It is clearly seen that the blood vessel is not in a stress-free state; rather, it varies enormously under constant cyclic pressure. This should be taken into account for stent development, in particular by studying the micro physical structure and micro mechanical properties of a chosen material theoretically and experimentally prior to stent development. The reason that micro structure and micro mechanics need to be addressed is because stent structure is at a micro scale, as shown in Figure 3. For Abbott’s ABSORB stent, which is made of PLA, the basic material structure is at even smaller scales of nanometres.

Figure 2: Hoop stress of a blood vessel as function of blood pressure and internal diameter (at fixed wall thickness of 0.8 mm)

Figure 3: SEM microstructure of a section of a stent.
(bar scale = 2 mm)
Figure 4 shows the basic chemical and physical structures of PLA. Taking into account the long range of amorphous and crystalline phases, its dimension is in the nanometre range. During the polymer degradation, both structures vary with time, as illustrated. The microstructure is not as simple as it looks. For example, the ratio of the two phases of amorphous and crystalline structures (crystals %), the size and size distribution of each crystalline and amorphous structure are all variables.

Figure 4: Basic chemical and physical structure of polyactide and variation of long range amorphous and crystalline structure during degradation (small angle X-ray scattering)
These variations will affect the mechanical properties of the PLA and, hence, the performance of the stent made from the polymer. The complexity is even worse during the polymer degradation, which always starts from the amorphous phase. The reason for the variation is because of the chiral nature of lactide (lactic acid), which lacks an internal plane of symmetry. In configuration there are two forms of lactide: D- and L. These two lead to two polymers poly(L-lactide), which is commonly used as a biodegradable polymer, and poly(D-lactide). Randomised polymerisation of the L- and D- form leads to PLA with no crystalline phase, that is, 100% amorphous phase.
There may be potential problems with a stent if it only uses PLA. This is because the glass transition temperature (Tg), which is a characteristic physical property of polymer, is approximately 60°C for poly(L-lactide) and the PLA in mixed ratio of L- and D- forms.6 This means that the PLA is a brittle material in applications at 37°C and room temperature. Because of a lack of elasticity (no rubbery phase in the PLA), brittle failure may possibly occur.

Choosing the best materials
Without doubt, biodegradable polymers are the choice for DEBS development because they have been used for surgical materials, including bone fixation plates, sutures, drug delivery devices, pharmaceutical applications7 and in scaffolds for tissue engineering applications.8

Biodegradable aliphatic polyesters, for example, polyglycolide (PGA), PLA and poly (ε-caprolactone) (PCL), represent a preferred choice of material due to their enhanced hydrophilicity and water solubility. A combination of PLA with PCL, with or without PGA, should make better DEBS than using PLA alone. For example, due to its Tg of just 60 °C as defined above, it is easy to damage PLA during processing or at a handling temperature of approximately 20°C. To overcome this mechanical property problem, PCL can be added to a PLA chain using chemical synthesis or mixed with the polymer to make a toughened stent. Improvements arise because PCL has a Tg of approximately −60°C.9 The amorphous phase of PCL acts as a rubbery material within the system so that the stent is no longer brittle. The same goal can also achieved using other biodegradable polymers where the Tg is well below room temperature.

Figure 5 lists the basic material structure for consideration during development of new DEBS using biodegradable polymers, including crystalline phase and amorphous phases. A combination of these three elements in the right concentration and structure should make a better DEBS than one made of PLA alone. However, there is scope for further development to achieve an ideal solution, for example, by employing drug-loaded glass to make a hybrid polymer composite. This can achieve two goals. First, to have better controlled drug release rate and second to further strengthen the mechanical property due to the contribution from glass particles in the system. Illustrative drug-loaded nano glasses are shown in Figure 5 as basic elements for the stent technology.
Figure 5: Basic "bricks” for building up the ideal DE
Further enhancement of DEBS can be achieved by developing hybrid systems and/or by employing drug-loaded soluble glass or sol-gel glass for better controlled release rate of the drugs. Through design it is best to use the four material characteristics stated above to make future DEBS satisfactory for their intended use.

Surface characterisation
There are contrasting needs in terms of the surface roughness associated with stent structures. Although roughness can present a problem in terms of damaging the intima during catheter deposition to the target site, it can also encourage endothelial cell deposition.

Use of techniques such as helium ion microscopy to etch nano-structures into surfaces represents a useful route to encouraging endothelialisation. A range of surface characterisation instruments allow chemical analysis from the top 2 to 10 nanometres downwards and high quality topographical data. This is particularly applicable for stent development and will be covered in the next article.

Future choices
Stent technology has been the latest innovation in percutaneous coronary intervention (PCI) with the advent of a CE-marked drug-eluting bioresorbable stent. This is a completely new technology and will be the future for PCI. The latest bioresorbable stent technology is just the beginning of a long journey with, perhaps, far more challenges than expected. The design and choices of materials will be the critical to success and this needs to be investigated thoroughly and systematically. This applies particularly to microstructure, microphysics and micromechanics as well as to the common practice of biological response studies, all of which must be considered prior to stent development and before design control commences.


1. Abbott Vascular, 11 January 2011.
2. J. Daemen et al., "Early and Late Coronary Stent Thrombosis of Sirolimus-Eluting and Paclitaxel-Eluting Stents in Routine Clinical Practice: Data from a Large Two-Institutional Cohort Study,” Lancet, 369, 667–78 (2007).
3. R. Waksman, "Biodegradable Stents: They Do Their Job and Disappear,” J. Invasive Cardiology, 18, 2, 70–74 (2006).
4. R. Fattori and T. Piva, "Drug-Eluting stents in Vascular Intervention,” Lancet, 381, 247– 249 (2003).
5. M.L.P.M. Verhoeven et al., "DSIMS Characterisation of a Drug-Containing Polymer-Coated Cardiovascular Stent,” J. of Controlled Release, 96 113– 121 (2004).
6. Yoshikazu et al., NATAS 2009, 37th Annual Conference, Lubbock, Texas, USA.
7. K.E. Uhrich et al., "ChemInform Abstract: Polymeric Systems for Controlled Drug Release,” Chem. Rev., 99, 3181–3198 (1999).
8. A. Atala, "Technology Insight: Applications of Tissue Engineering and Biological Substitutes in Urology, Nature Clinical Practice Urology, 2, 143–149 (2005).
9. Y. Tokiwa et al., "Biodegradability of Plastics,” Int. J. Mol. Sci., 10, 3722–3742 (2009).
    Dr Xiang Zhang is Principal Consultant, Medical Materials, and
Phil Jackson is Business Development Manager, Medical Materials
both at CERAM, Queens Road, Penkhull,
 Stoke-on-Trent ST4 7LQ, UK,
Tel. +44 (0)1782 764 428,


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