Still Encapsulating Your Stent in Two ePTFE Layers? One Layer. Bonded.

Why the two-layer default deserves a second look

Many ePTFE covered stents are still built the way they have been for decades: the metal frame is sandwiched between an inner and an outer ePTFE layer, and the two layers are fused to each other through the stent cells with heat and pressure. The method is proven and clinically familiar. It also locks four compromises into every device: a double wall, an unbonded pocket around each strut, a membrane whose pores are reshaped during fusion, and a build that depends on skilled manual work.

Sutureless lamination removes the sandwich. A single ePTFE membrane is formed first, under controlled conditions, then bonded directly to the struts with a thin elastic polymer binder. The sections below take the four differences in turn and set each one against the published evidence.

How two-layer encapsulation works

The construction is set out in patents first filed in 1995. An inner ePTFE tube goes on a mandrel, the stent is loaded over it, and an outer ePTFE tube is pulled over the stent. The assembly is wrapped under tension with PTFE tape, which “compresses the ePTFE/stent/ePTFE composite structure against the mandrel” so the two layers meet through the stent cells [1]. It then goes into an oven at or above PTFE’s 327 °C melting point [2]: the patent examples use 337 to 375 °C [1, 3], and a 2023 paper on ePTFE stent coverings gives 340 to 365 °C for sintering ePTFE [4]. A related route bonds an ePTFE lining to a nitinol frame with an FEP fluoropolymer adhesive at 360 to 380 °C [5].

The bond that forms is PTFE to PTFE, not PTFE to metal. In the original patent’s words, the layers bond “through the openings between adjacent struts, without the need for intervening adhesives, by mechanical interaction between the node-fibril microstructures” [3]. The strut itself is captured, not bonded. That is by necessity: PTFE wets and bonds poorly, with a critical surface tension of only 17.5 to 20.5 mN/m [6].

The result works and is clinically familiar, but it carries four built-in costs: two walls instead of one, a pocket around each strut, a membrane re-processed by heat and pressure on the device, and a bond whose quality is proven after the oven rather than before it.

1. Single layer, not two: a thinner wall and a lower profile

Every layer of cover adds to the crimped diameter. Typical ePTFE covering membranes are 100 to 200 µm thick [7], and an encapsulated device carries two of them. By Medibrane’s figures, a sintered two-layer wall typically finishes near 225 µm, while a single laminated ePTFE layer finishes at 20 to 30 µm.

That difference shows up in the delivery system. The authors of a randomized trial of covered versus bare stents noted that, because of the bulk of the additional graft material, “covered stents require larger delivery systems” [8]. A cover also raises the force needed to crimp a stent by more than 136% at full crimp [9]. Coronary covered stents offer a direct comparison of architectures: a design that sandwiches PTFE between two metal stents has a crossing profile of 1.63 to 1.73 mm [10], while a single-layer PTFE design reaches 1.1 to 1.4 mm and fits a 5 F guide catheter in every size [11].

Delivery size has clinical consequences. In a prospective registry of 5,385 peripheral endovascular procedures, access-site complications occurred in 16.6%, and each additional French of sheath size raised the odds by 59% (OR 1.59) [12]. A registry of 27,048 peripheral interventions identified sheath size above 6 F as a predictor of access-site complications [13], and in transfemoral TAVR, major vascular complications fell from 10.5% with larger sheaths to 0.5% with lower-profile sheaths [14]. The link is not universal [15], but it matters most where vessels are small: in 2,084 pediatric catheterizations, femoral artery occlusion occurred in 2.3% of cases when the sheath was large for the child’s age, against 0.4% otherwise [16].

2. Bonded to the struts: no voids where thrombus can start

When two layers close around a round strut, they cannot follow its contour all the way to the point where they meet. A small unbonded space remains at each strut shoulder, and the strut is held only by the pocket around it. Crimping, deployment and roughly 40 million cardiac cycles a year all work on that pocket.

Layer separation is a documented hazard. An FDA-commissioned review of PTFE device safety lists “graft and stent component separation” among the more serious hazards reported [17]. In multilayer ePTFE dialysis grafts, a case report described separated graft layers with thrombus in the false lumen between them [18], and a clinical series reported delamination in 5.1% of grafts by 21 months [19]. These are needle-punctured access grafts rather than stents, but they show what happens once blood reaches the space between layers.

Flow at the strut is the second concern. In a study combining bench flow models, animal testing and computational modeling, 162 µm struts were about 1.5 times more thrombogenic than 81 µm struts, and predicted flow disturbances “correlated with spatial distribution of formed clots” [20]. Recirculation zones beside struts act as “micro-reaction chambers where procoagulant and pro-inflammatory elements from the blood and vessel wall accumulate,” and a streamlined strut profile can produce “a local flow field free of recirculation zones” [21]. Gaps matter too: in an in vitro model, malapposed stent sections carried 13.9 mm² of thrombus against 0.41 mm² for well-apposed sections [22].

Lamination addresses both mechanisms. The membrane is bonded along each strut by an elastic polymer binder, so there is no pocket to fill or to separate, and the binder forms a smooth fillet where strut meets membrane. The bond also carries cyclic load in shear, the direction in which adhesive joints are strongest; peel strength can be as low as 1% of shear strength [23].

No head-to-head study has yet compared thrombus on laminated and encapsulated covers. The case rests on these established mechanisms, and every device design still needs its own bench and animal testing.

3. Controlled microstructure: make the membrane first, then attach it

ePTFE performance lives in its node-fibril microstructure. That structure is set during extrusion, stretching and sintering, and the stretch ratio, temperature and speed control fibril length (the internodal distance, or IND), porosity and thickness [2, 24]. Sintering also costs elasticity: in one industrial-scale study, elongation at break fell from 430% (transverse) and 189% (longitudinal) to about 75% as sintering temperature rose to 400 °C [24].

The body responds to that structure. In baboons, 60 µm IND grafts were covered with endothelium within weeks, fed by capillaries growing through the wall, while the central portions of 30 µm grafts “were still not healed at 3 months” [25, 26]. Golden and colleagues found “an abrupt switch in the pattern of healing as porosity was increased from 30 to 60 μm” [27]. In humans, endothelium grows only a short way in from the anastomosis, so ingrowth through the wall is the realistic route to a healed surface [28, 29]. Pore structure is therefore a design choice, made per indication.

Encapsulation re-processes that structure on the device. Wrap tension presses the layers together through the stent [1], and because the struts stand proud of the mandrel, the membrane over each strut takes more of that pressure than the membrane between them. Heat-bonding ePTFE layers around a stent produced “a significant decrease of the pores among the fibers” and “obvious fusion” of fibers [30]. Reheating finished ePTFE to 350 °C reduced its porosity and stiffened it by 25 to 73% [31], and even modest pressure of 1 to 45 kPa reduces the porosity and thickness of PTFE membranes [32]. Strut-level densification has not been mapped in a published study, but these findings all point the same way.

Lamination reverses the order. The membrane is produced on its own mandrel and characterized first, so thickness, IND and porosity are measured and accepted before it meets the stent. Medibrane then bonds it at roughly 100 to 200 °C with a thin elastic binder, far below PTFE’s melting point, so the accepted microstructure is the one that goes into the device.

The lower temperature also protects the frame. Nitinol transformation temperatures shift when the alloy is aged at 300 to 500 °C, with the fastest precipitation near 425 °C [33], and FDA guidance notes that they are “sensitive to small variations in manufacturing processes, such as temperature or heating duration” [34]. A sintering cycle at 337 to 380 °C sits inside that band. A lamination step at 100 to 200 °C sits well below it.

4. A semi-automatic process: high yield, repeatable

Covered-stent performance depends on tolerances. In finite element models, cover thickness produced “a non-linear increase in radial stiffness” and in chronic outward force [35]. Deployment force, the force the physician applies to release the device, is driven by radial force and friction, and both are set by the stent and the cover together. Variation in either one shows up as scatter in deployment force, part to part and lot to lot, which is why it has to be tested and controlled.

Encapsulation makes that control harder. The build runs as a sequence of layup, wrapping, oven and trim steps [30], and the quality of the sintered bond cannot be fully confirmed by inspecting the finished part. Regulators treat such steps as special processes. The former 21 CFR 820.75 required that where results “cannot be fully verified by subsequent inspection and test, the process shall be validated with a high degree of assurance” [36], and the same principle carries into FDA’s Quality Management System Regulation, effective February 2, 2026, which incorporates ISO 13485:2016 [37]. GHTF guidance calls validation “vitally important if the predetermined requirements of the product can only be assured by destructive testing” [38]. Visual inspection is a weak backstop: inspectors of precision parts correctly rejected 85% of defective items and wrongly rejected 35% of acceptable ones [39].

Lamination is designed for process control. Because the membrane is inspected and accepted before bonding, defects are caught upstream rather than after the oven. Medibrane runs the bonding step semi-automatically with in-process checks, and the low process temperature allows a part to be reworked rather than scrapped. The aim is fewer rejects and tighter part-to-part and lot-to-lot consistency, which is what keeps deployment force predictable in the physician’s hand.

What this means for device teams

The construction choice sets profile, strut interface and process window early in a program, so it is worth making deliberately.

Attribute	Two-layer encapsulation	Sutureless lamination
Cover wall	Two ePTFE layers, about 100 µm typical	One ePTFE layer, 20 to 30 µm
Strut interface	Captured in an unbonded pocket	Bonded along the strut by an elastic binder
Process temperature	337 to 380 °C	About 100 to 200 °C
Membrane microstructure	Re-processed by heat and wrap pressure on the device	Made and accepted first, then attached
Quality evidence	Proven after the oven, often destructively	In-process checks before and during bonding
Cover options	Full sandwich on both surfaces	Inner or outer surface, full or selective zones

Teams working toward a smaller delivery system, a nitinol frame with tight transformation targets, or a membrane with a specified pore structure should evaluate lamination at the concept stage, before the design freezes around the sandwich.

Medibrane develops and manufactures sutureless-laminated ePTFE covers for stents and implants as an FDA-registered, ISO 13485:2016 certified CDMO. To discuss your device, book a 20-minute consultation.

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