A compressed stent self-expands to full diameter once warmed by the body
Before it ever reaches a patient, the stent is heated above its transition temperature, formed into an expanded lattice cylinder, then cooled while compressed onto the catheter shaft — mechanically locking in a temporary, low-diameter shape.
The stent is fabricated at full diameter, heated above Ttrans to become rubbery, mechanically compressed onto the catheter, and cooled below Ttrans while held in compression. The polymer chains freeze into this strained, temporary configuration — a process called shape programming.
Biodegradable shape-memory polymers are typically PLA/PCL (polylactic acid / polycaprolactone) copolymer networks. Crystalline or glassy "switching" segments lock the temporary shape, while covalent or physical netpoints "remember" the original expanded geometry.
Because Ttrans sits close to but below 37°C, the compressed shape is stable at room temperature and during cold-chain storage — the stent will not prematurely deploy before it reaches the patient.
Unlike Nitinol self-expanding stents, which rely on a temperature-driven martensite–austenite phase transformation in a metal alloy, SMP stents store and release strain energy through polymer chain network mechanics — a fundamentally different physical mechanism that also enables full biodegradability.
The compressed stent, still locked below its transition temperature, is advanced through a catheter sheath along the vascular or airway pathway toward the lesion — its small profile minimizing trauma during transit.
A small crimped profile allows the delivery system to navigate tortuous, narrowed, or pediatric-scale vessels and airways that would not accommodate a bulkier balloon-expandable metal platform.
Ambient body warmth near the catheter shaft is normally insufficient to trigger premature expansion, since the polymer is engineered with a sharp transition just under 37°C and the compressed segment is largely shielded within the sheath.
Radiopaque markers embedded in the stent or catheter tip let the interventionalist track position under fluoroscopy or endoscopy throughout the delivery pathway.
With the catheter tip aligned across the stenosis or collapsing segment, the outer restraining sheath is withdrawn, exposing the compressed stent while it is still below its transition temperature.
Retracting the outer sheath removes the mechanical constraint holding the stent compressed. Blood or airway humidity at close to body temperature begins conducting heat into the polymer struts.
Because expansion is not instantaneous, clinicians have a short working window to fine-tune stent position before the lattice locks into its final expanded geometry — an advantage over instantly-deploying platforms.
A small amount of shape recovery begins immediately as the outermost struts contact warm blood, gently seating the stent against the vessel wall before full thermal equilibrium is reached.
As the polymer equilibrates to 37°C body temperature, it crosses its transition temperature and softens; the elastic strain stored during programming drives the lattice to recover its original expanded diameter, gently but firmly opening the vessel or airway.
Unlike balloon-expandable metal stents, which need mechanical over-pressure from an inflated balloon to plastically deform the metal, the SMP stent expands from its own stored elastic energy once warmed — eliminating a step of the procedure and the associated barotrauma risk.
Recovery ratio describes how close the deployed diameter comes to the fully programmed target shape. High-fidelity SMP formulations reliably reach 85–98% recovery in vivo, with the remainder recoverable creep occurring over the following hours.
Radial force — the outward pressure exerted against the vessel or airway wall — builds progressively as the lattice unfolds, reaching a therapeutic plateau sufficient to resist elastic recoil of the diseased tissue without overstressing the wall.
Each platform trades off delivery mechanism, permanence, and radial strength differently — SMP stents are distinguished by combining a low-profile, gentle thermal deployment with true biodegradability.
At full radial force the stent holds the vessel or airway patent through the healing phase. Over the following months, hydrolytic degradation of the PLA/PCL backbone gradually weakens and thins the struts as the surrounding tissue takes over mechanical support.
The degradation timeline is tuned via polymer molecular weight, crystallinity, and copolymer ratio so that mechanical scaffolding persists only as long as the vessel or airway needs it — typically 6 to 24 months depending on the clinical indication.
Ester bonds in the PLA/PCL backbone are cleaved by water, progressively lowering molecular weight, then strut cross-section, until the lattice loses mechanical integrity and is cleared as lactic acid and other resorbable byproducts.
Because no permanent metal cage remains, the risks associated with lifelong foreign-body presence — late stent thrombosis, chronic inflammation, restricted future vessel growth, and the need for a second removal procedure — are avoided entirely.
This is especially valuable in pediatric vascular and airway stenting, where a permanent, non-degradable implant like Nitinol cannot expand with a growing child and may eventually require surgical removal or replacement.