Feature Article

As medical technologies evolve and the performance of biomaterials within the human body becomes better understood, there appears to be an increasing demand for degradable and resorbable polymers. New strategies are being developed to meet this demand.

By: David Williams

For many years a principal objective of polymer scientists in the biomaterials field was to select or design materials that would resist degradation within the human body. Indeed many of the problem areas in biomaterials science have been related to the difficulty of achieving sufficient inertness in the chosen polymers. The stress cracking of polyurethanes in pacemaker leads, the oxidation of polyethylene in joint prostheses, the slow hydrolysis of aromatic polyesters in vascular grafts and the degradation of silicone elastomers in lipid environments have all led to clinical failures and stimulated the development of even more degradation-resistant polymers. In many areas of medical technology, however, the situation is being reversed. More applications require that the polymer degrades and the objective of many of today’s biomedical polymer scientists and engineers is to design polymers that degrade with precisely the correct characteristics that meet the specifications of new devices. These specifications can be summarised by the need to degrade with the appropriate kinetic profile and to minimise, or preferably eliminate, the potential for degradation products to do harm, while at the same time achieving the intended functional performance of the device, which may well be pharmacological or biological. These new objectives may require a new approach to polymer design, or at the very least, a much better understanding of the mechanisms of polymer biodegradation within the environment of the human body.

 

Let us take one example of a medical device where this change in concept is currently taking place. This involves intravascular stents. These have been in clinical use now for more than 20 years and probably represent one of the most widely used implantable devices. They have been associated with a complex interplay between engineering design and metallurgical structure so that optimal devices can be used that display exactly the right characteristics of the elastic deformation required for endovascular insertion of the stent and the plastic deformation required for deployment. The result has been a group of commercially available stents with good engineering qualities. The problem, as everyone knows, is that these bare metal stents do not, and probably cannot, prevent restenosis because they continue to irritate the endothelium of the artery for as long as they remain in place. One solution has been to incorporate powerful anti-proliferative drugs such as the chemotherapeutic agents paclitaxel and sirolimus into the surface layer of the stents, which can minimise the harmful hyperplastic response of the endothelium and lead to in-stent restenosis. Although it is possible for these drugs to be attached to the metal, most attention has been paid to their incorporation into a thin layer of a polymer surface coating. Commercially available drug-eluting stents have all opted for nondegradable coatings at this stage and sirolimus-poly-n-butyl methacrylate based copolymer and paclitaxel-polystyrene-b-isobutyl-b-styrene combinations are among the leaders here. The drug is released over a matter of weeks and is usually effective in controlling hyperplasia, but this relief is not guaranteed to last forever. There is also a small, but finite, risk of serious thrombo-embolic events occurring after the drug has been eluted. The essential problem remains: as long as there is an irritating object within the lumen of the artery there will be the possibility of further endothelial proliferation. The answer some authorities argue is to use a totally degradable polymeric stent.¹ In passing it should be noted that others argue that a degradable magnesium alloy could also be used. It is not clear yet whether the deformation characteristics of the metal systems can be replicated for this application or whether the biocompatibility of the degrading system within the vasculature is sufficiently good, but the early signs are impressive. The first results, published recently, from a clinical trial of a everolimus-oly(lactic acid) fully absorbable stent look good.² It may well be that the biomaterials used in this pivotal medical device will totally change.

 

The degradation of a polymer can take place through a number of different mechanisms, but the end result is essentially the same: the long chains that constitute the polymeric structure are broken down into smaller, usually very much smaller, units. If the intention of using a degradable polymer in a medical application is to allow a device to perform some function on a temporary basis, as with the stent or bone fracture plates and other products of this type, then the specifications should require that the polymer has the essential characteristics that allow it to perform that function over the required period of time. It should then break down into small units at a physiologically relevant rate, preferably uniformly, without the small units having any adverse effect on the host, locally or systemically. Bearing in mind that many polymers contain a wide variety of additives and residual catalysts, great care has also to be taken to avoid any harmful effects associated with the eventual release of these substances. If the function of the polymeric material is biological, as with a tissue engineering scaffold where the material surface should control cell behaviour, or if it is pharmacological, as with drug or gene delivery systems, then the degradation process should enhance those activities rather than antagonise them.

 

Critical to the satisfactory compliance with these specifications are the facts that most monomers or low-molecular-weight derivatives of polymers are toxic or irritant and most degradation processes are heterogeneous, that is, the breakdown occurs irregularly. This often leads to intermediate situations where fragments of varying size are released from the polymer before the final end products are formed; these fragments are themselves usually irritants. The situation does not, in general, look good as far as achieving optimal biocompatibility and biological safety with degradable systems are concerned.

 

We cannot, therefore, expect to design suitable degradable polymers for medical applications around the vast majority of commercially available materials. The best strategies for many situations involve the design of polymers that can uniformly break down in the body by a simple, predictable mechanism that is independent of where in the body the device is located and of the precise nature of the tissue environment. The simplest mechanism is that of hydrolysis, which here means the interaction between the polymer and water. The water cleaves bonds between atoms in the backbone of the molecular chain such as in ester groups. This requires the water to gain access to the molecules in a predictable way, for example, by diffusion into the polymer bulk or surface erosion. It is important that this process is not unduly influenced by the biological activity of the tissue, for example, through the effects of enzymes or reactive oxygen species released from cells, because this increases unpredictability of the degradation rate and the morphology of the intermediate products. It is fortunate that there are sufficient polymeric systems that do degrade by this type of hydrolysis to give us a number of potential degradable polymeric platforms. Incidentally, this hydrolysis is really a degradation process rather than a biodegradation process, although with respect to the human body those terms tend to be used interchangeably.

 

If we assume that we have a uniform process of hydrolysis, we now have to consider the nature of the end products. Because it is usually our intention that the degradation products are eliminated from the body, we have to make sure that these products, which could be monomers, oligomers or other small molecules, do not irritate the tissue or have harmful effects during elimination, which would often be via the liver or other organs. The best way to minimise this risk is to ensure that these end products are easily metabolised in the tissue rather than eliminated by a detoxification process in the liver or even worse, stored somewhere in the body in a nonphysiological form, where effects such as genotoxicity could arise. The best products are those that are metabolised into simple molecules such as water and carbon dioxide during natural tissue processes. They are therefore resorbed, or absorbed, by the tissue and leave no trace. This is why we seek degradable, resorbable polymers for these applications. The polymers and copolymers based on lactides and glycolides and those based on certain proteins and polysaccharides come into this category, as do a number of others. 

 

The future of these systems looks more promising as they become refined and adapted to the diverse applications in medicine. Some specific examples of these newer degradable resorbable polymers will be given in a later column. 

 

Morgan & Masterson, Avenue de la Forêt 103, Brussels 1000, Belgium tel. +32 4 7597 0556 
e-mail: peggy@morgan-masterson.com