After the drug is gone, the polymer disappears too — reducing the long-term foreign-body burden
A drug-eluting stent (DES) is a metal scaffold coated with a thin polymer layer that carries and releases an antiproliferative drug to prevent restenosis — the re-narrowing of a treated artery. In bioresorbable-polymer platforms, that coating is not meant to stay forever: it is engineered from hydrolytically degradable polyesters such as poly-L-lactic acid (PLLA) or poly(lactic-co-glycolic acid) (PLGA) that dissolve on their own once their job is done.
Bare metal stents alone provoke smooth-muscle cell proliferation and neointimal overgrowth that can re-narrow the vessel (restenosis) in 20–30% of cases. Drug-eluting stents solved this by loading an antiproliferative drug (sirolimus, paclitaxel, everolimus, zotarolimus) into a polymer carrier bonded to the strut surface. The polymer performs two jobs: it controls the rate and duration of drug release, and it physically anchors the drug to the metal so it is not washed away before reaching the tissue.
Early-generation DES (Cypher, Taxus, ~2002–2006) used permanent, non-degradable polymers such as PEVA/PBMA or SIBS. These worked well for suppressing restenosis, but once the drug payload was fully released — typically within 30–90 days — the polymer had no further purpose. It simply remained on the strut indefinitely as a permanent foreign material.
A polymer left behind after its drug is spent is not inert forever. Durable polymer coatings on first- and second-generation DES were increasingly linked to chronic low-grade vascular inflammation, delayed and incomplete strut endothelialization, and hypersensitivity reactions in a subset of patients — all of which are mechanistically tied to very late stent thrombosis (VLST), a rare but often fatal complication occurring more than one year after implantation.
Bioresorbable-polymer DES were designed on a simple engineering premise: give the polymer a finite service life matched to the drug-release window (weeks to a few months), then let it hydrolyze away entirely over the following months, leaving nothing but the metal scaffold and — ideally — a mature, healed vessel wall. This class of device, alongside polymer-free DES, represents the industry's direct response to the durable-polymer VLST signal.
In pooled meta-analyses, durable-polymer first-generation DES showed very late stent thrombosis rates continuing to accrue at roughly 0.4–0.6% per year beyond year one, whereas newer bioresorbable-polymer and polymer-free platforms have shown a plateau in late thrombotic events after the coating resorbs — the central clinical motivation for this technology.
The degradation clock starts the moment the coating is exposed to an aqueous, protein-rich physiological environment. Water molecules diffuse into the polymer matrix and begin attacking the ester linkages that hold the polyester backbone together — a purely chemical process, hydrolysis, that requires no cells or enzymes to proceed.
PLLA and PLGA are aliphatic polyesters: their backbones are built from repeating lactic-acid (and, for PLGA, glycolic-acid) units linked by ester bonds. Water molecules that diffuse into the matrix attack the carbonyl carbon of these ester linkages, splitting the chain into two shorter chains and generating a new carboxylic acid end group and a new hydroxyl end group at the break point.
This reaction is autocatalytic: the carboxylic acid groups produced by each cleavage event locally lower the pH and catalyze further hydrolysis nearby. Because water can penetrate the entire amorphous fraction of the polymer relatively quickly, cleavage events begin distributed throughout the material's volume — not just at its outer surface.
The lactide-to-glycolide ratio (for PLGA) or the use of pure PLLA sets the degradation timeline. Glycolide units are more hydrophilic and less sterically hindered, so PLGA copolymers with higher glycolide content hydrolyze faster than lactide-rich or pure PLLA polymers. A 50:50 PLGA can lose the bulk of its molecular weight within weeks to a couple of months, while a slower, more crystalline PLLA homopolymer can take well over a year to reach the same point — the basis for the "Polymer Type" comparison in this simulation.
At this stage, molecular weight is already falling measurably, but the coating's outward mass and physical shape remain essentially unchanged — the defining signature of bulk erosion, which is explored further in the next stage.
Because early molecular-weight loss produces no visible change in coating mass or stent appearance, hydrolysis at this stage is often described as "invisible" degradation — detectable only by gel permeation chromatography (GPC), not by eye or by simple gravimetric weighing.
Hydrolysis continues throughout the entire thickness of the coating, not just at its exposed surface — a mechanism called bulk erosion. Long polymer chains are progressively cut into shorter oligomers, and molecular weight falls by an order of magnitude, even while the coating still looks and behaves like a solid, continuous layer.
Degradable polymers fall into two broad erosion classes. In surface erosion (typical of polyanhydrides and some polyorthoesters), water cannot penetrate the bulk quickly, so hydrolysis is confined to a thin outer shell that erodes away layer by layer — the object shrinks but its interior stays at full molecular weight until the very end.
PLLA and PLGA behave oppositely: water diffuses into the entire matrix faster than hydrolysis consumes it, so chain scission happens simultaneously throughout the coating's volume. The practical consequence is the molecular-weight-versus-mass-loss distinction that defines this stage: Mw can fall by 80–90% while the polymer has lost only a small fraction of its physical mass, because the fragments are still too large to dissolve and diffuse out of the matrix.
As carboxylic-acid chain-end groups accumulate inside the coating, they cannot easily diffuse out of a still largely intact matrix, so acidity builds up preferentially in the interior relative to the surface (which can buffer against the surrounding tissue fluid). This creates a self-accelerating "autocatalytic core" effect — the center of a bioresorbable coating or scaffold strut often degrades faster than its outer rim, sometimes leading to a hollowed-out or honeycombed internal structure well before outward dimensions change appreciably.
Mechanically, this stage is also where the coating quietly loses much of its tensile and shear strength, even though it has not yet begun to visibly fragment — a gap between mechanical failure and visible mass loss that matters for how long a bioresorbable scaffold structure (not just a coating) can be relied upon for radial support.
Gel permeation chromatography studies of PLGA coatings typically show molecular weight falling to 10–20% of baseline within the first several months, while gravimetric mass loss over that same window is often under 10% — a striking illustration of bulk erosion's decoupling of Mw decline from mass loss.
Once oligomer fragments shrink below roughly 1,000–5,000 Da, they finally become small enough to dissolve in tissue fluid and diffuse out of the coating. Physical mass loss — the visible thinning and fragmentation of the coating — now proceeds rapidly, and the released fragments are cleared through completely normal, non-toxic metabolic pathways.
As chain scission continues, an increasing fraction of the coating consists of oligomers and monomers small enough to be water-soluble. These diffuse out of the remaining solid matrix into the surrounding tissue fluid, while larger microparticle fragments that break off mechanically are engulfed by local macrophages and giant cells via phagocytosis — the same foreign-body-clearance machinery the body uses for other resorbable implants.
The end products of PLLA/PLGA hydrolysis are lactic acid and glycolic acid — small, physiologically ubiquitous molecules. These are metabolized through completely ordinary cellular pathways: lactic acid is oxidized via the Krebs (citric acid) cycle, and glycolic acid is either oxidized similarly or excreted renally. The terminal products are carbon dioxide and water, exhaled and excreted through completely normal physiology. This is the same degradation chemistry that has been used safely in resorbable surgical sutures for over 40 years.
Because bulk erosion keeps mass roughly constant for a long induction period while Mw silently collapses, mass loss can appear to happen quite abruptly once the "critical molecular weight" threshold is crossed throughout the coating simultaneously. In this stage, the coating visibly thins, develops cracks and voids, and begins shedding small resorbable fragments — a dynamic that under polarized light or SEM imaging looks like a sponge-like breakdown rather than a slow, uniform shrinkage.
This is also the stage where local tissue response is most active: transient mild inflammation associated with fragment phagocytosis is expected and typically resolves as clearance completes, in contrast to the chronic, low-grade inflammation associated with permanent durable-polymer coatings that never fully clear.
The lactic-acid and glycolic-acid monomers released during clearance are chemically identical to metabolites the body produces routinely during normal glycolysis and anaerobic exercise — there is no exotic or foreign chemistry involved in disposing of a fully hydrolyzed PLLA/PLGA coating.
Roughly 12–24+ months after implantation, depending on the specific polymer chemistry and coating thickness, the bioresorbable coating has fully hydrolyzed, fragmented, and cleared. What remains is the bare metal stent strut, by this point ideally embedded beneath a mature, stable neointimal and endothelial layer — the vessel's own healed lining covering the scaffold from the inside.
Randomized trials and large registries have compared three DES coating strategies: durable-polymer (permanent, e.g. everolimus-eluting Xience), bioresorbable-polymer (e.g. Synergy, Ultimaster, Orsiro), and polymer-free (drug bonded directly to a micro-structured metal surface, e.g. BioFreedom). Bioresorbable-polymer platforms have generally demonstrated non-inferior restenosis and target-lesion revascularization rates compared with best-in-class durable-polymer DES, alongside numerically low and often plateauing very late stent thrombosis rates once the coating has resorbed.
Polymer-free DES pursue the same goal — eliminating chronic polymer exposure — by never applying a bulk polymer layer at all, instead relying on a porous or textured metal surface to hold the drug transiently. Across these platforms, the shared clinical objective is the same: minimize the duration during which a foreign, potentially pro-inflammatory material remains in contact with the arterial wall after the drug's therapeutic window has closed.
If a coating can safely disappear, the logical extension is a scaffold whose entire structural backbone also resorbs — a fully bioresorbable vascular scaffold (BVS), leaving no permanent metal behind at all. The best-known example, Absorb (everolimus-eluting PLLA scaffold, Abbott), was designed to provide temporary radial support before resorbing over roughly two to three years, restoring the vessel's natural vasomotion and leaving nothing implanted.
However, thicker BVS struts (needed for adequate mechanical strength from PLLA alone) were linked to higher rates of scaffold thrombosis compared with contemporary metallic DES in trials such as ABSORB III, leading to Absorb's market withdrawal in 2017. This outcome reinforced an important lesson directly relevant to coating design: degradable polymer engineering must balance mechanical performance, degradation kinetics, and thrombogenicity together — not resorbability alone. Newer-generation scaffolds, including thinner-strut polymer designs and resorbable magnesium-alloy scaffolds, continue to pursue this goal.
The evolution from permanent-polymer coating → bioresorbable-polymer coating → polymer-free DES → fully bioresorbable scaffold traces a single engineering thread: progressively minimizing the amount of permanent foreign material left in the artery after healing is complete, while keeping the acute mechanical and antiproliferative benefits that make stenting effective in the first place.