Bioresorbable stent scaffolds are balloon-expandable and have been used to replace metallic stents to treat the narrowing of arteries and airway passages. Like traditional metallic scaffolds, bioresorbable scaffolds provide artery and/or airway support and act as a delivery system for the controlled release of an anti-inflammatory drug to prevent narrowing of the vessel lumen. Unlike metallic scaffolds that are permanent, bioresorbable scaffolds degrade over time and are completely absorbed and eliminated from the body after they are implanted, and therefore eliminate the need for additional surgical procedures. Bioresorbable scaffolds are used in coronary, peripheral, neural, and nasal applications. Each application has different scaffold sizes (OD and wall thicknesses), vessel support requirements, and degradation times.

Bioresorbable Polymers

Fig. 1 – Shown is an example of scaffold vessel support.

Bioresorbable polymers, such as poly-L-lactide (PLLA) and polylactic-co-glycolic acid (PLGA), are used for stent scaffolds and other medical applications such as resorbable sutures, ligament tissue, orthopedic implants, and controlled drug delivery systems. Both PLLA and PLGA degrade by hydrolysis, which is the chemical breakdown of the compound due to the reaction with water. The glass transition temperature (Tg) of both PLLA and PLGA is higher than the normal body temperature of 37°C, which allows the polymers to remain rigid and resistant to short-term creep deformation after implantation.

PLLA and PLGA are high molecular weight polymers. For example, a common PLLA used for a coronary scaffold has inherent viscosity (IV) of 3.8 dl/g compared to PET, which may have an IV of .75 dl/g. PLLA is a semi-crystalline biocompatible homopolymer supplied in granule form that has high mechanical strength ideal for extruded bioresorbable coronary stent scaffolds. Typical PLLA used for scaffolds has a melt temperature (Tm) of 180° to 190°C (356° to 374°F) and is heated to 200° to 210°C (390° to 410°F).

PLGA is a biocompatible copolymer that is commonly used for nasal scaffold implants and suture applications. PLGA devices degrade faster than their corresponding PLLA devices, which also have greater mechanical properties.

In addition to PLLA and PLGA bioresorbable polymers for scaffolds, the surface of the scaffold is coated with a bioresorbable polymer and a drug that is released over time to prevent restenosis (narrowing of the stent lumen). Poly-DL-lactic acid (PDLLA), which is an amorphous (non-crystalline) polymer, is used as a carrier of controlled-release anti-inflammatory drugs, such as Everolimus or Sirolimus, that prevent restenosis.

Manufacturing Process

Bioresorbable stent scaffolds are manufactured from PLLA and PLGA polymers that are extruded into tight tolerance tubing that is secondarily radially and axially expanded, and laser machined into a scaffold pattern. Additional processes include surface coating the scaffold with a bioresorbable polymer that carries a drug, adding radio-opaque marker bands on each end, crimping the stent to the balloon catheter delivery system, and radiation sterilization.

Processing bioresorbable materials for implants is typically done in a cleanroom environment with current Good Manufacturing Practices (cGMP). cGMP requirements involve the maximum use of stainless steel, special material selections, and all surfaces that come in contact with the product must be smooth and without any raised or recessed points where materials can be trapped and degrade. The cGMP requirements for processing bioresorbable materials are typically not as stringent as the requirements for pharma applications.

Mechanical Properties and Molecular Weight

Critical to the functionality of a PLLA scaffold are certain mechanical properties, mainly a high radial strength for vessel support and to resist inward recoil of the scaffold after implantation within a desired time. (See Figure 1) The scaffold must also be flexible enough to be delivered distally into the artery via balloon expansion without fracturing the struts of the scaffold. PLLA is brittle and scaffolds have been known to have strut fracture due to over-dilation during balloon expansion. For this reason, PLLA copolymers have been developed that are strong but more ductile to prevent fracture during excessive balloon expansion.

The mechanical properties of the PLLA and PLGA scaffolds are greatly reduced by thermal degradation of the polymer and molecular weight loss caused from excessive thermal history (temperature and time) during the extrusion process. Moisture content prior to extrusion and secondary exposure to e-beam or gamma radiation sterilization also has an impact on molecular weight loss. The molecular weight (Mn) of polymers determines the mechanical properties of polymers and it is measured in Daltons (Da). The Mn is greatly reduced during the manufacturing process by more than 70 percent, depending on the moisture content prior to extrusion, the thermal history during extrusion, and the radiation dose applied after extrusion.

Keys to Optimizing Extruded Scaffold Tubing

The higher the thermal history during extrusion of the tubing, the greater the decrease in Mn, and decrease in the time that the radial strength of the scaffold is effective. Therefore it is critical to minimize the residence time and process temperatures during extrusion of bioresorbable resins to maximize absorption time and the radial strength of the extruded PLLA tubing. PLLA has often been processed on 19mm (¾") and 25mm (1") extruders but this is not recommended due to thermal degradation caused by excessive residence time with an over-sized extruder. For future scaffold developments, PLLA should ideally be extruded on micro extruders in the 12mm to 16mm screw size range. Extruding a 2.5mm (0.098") OD x 150μm (0.006") wall thickness tube from a PLLA resin on a 19mm (3/4") extruder at 5-10 RPM screw speed will result in a residence time of 8 to 10 min. Extruding the same PLLA tube on a 12mm micro extruder at 15 to 20 RPM screw speed will result in a residence time of 5 minutes or less.

Special attention is paid to the proper motor sizing and/or screw speed setup to avoid utilizing elevated barrel temperature settings. The low processing temperatures of bioresorbable polymers increase the torque applied to the feed screw, which could result in screw breakage.

Excessive moisture during extrusion reduces the molecular weight and has a negative impact on the final stent performance. Therefore it is recommended that PLLA and PGLA granules are properly dried prior to extrusion with a nitrogen dryer system or desiccant dryer that can achieve below -40°F dewpoints so the resin can be dried down to a very low moisture level of 50 ppm. To keep resin dry at the feed throat, it is also recommended to introduce a blanket of nitrogen into the transition piece between the machine-mounted drying hopper and extruder feed throat. This is critical for applications where the resin is outside the drying environment for more than 15 minutes.

Fig. 2 – Micro extruder with polished stainless shroud for cGMP.

Proper cleaning and maintenance of the extrusion machine in a cGMP environment will help ensure that residues and contamination will not be extruded into the scaffold tubing. This means the extruder should be designed in such a way that allows the screw, barrel, and feed insert of the extruder to be easily removed for unrestricted bare metal cleanouts. In addition, cGMP involves the maximum use of polished stainless steel shrouds with rounded corners to protect product contact surface areas from contamination and for easy cleaning. Figure 2 shows a 12mm micro extruder designed to meet cGMP requirements.

Imperfections, such as gels, within the polymer melt come from material inconsistencies, degradation, and contamination and should be prevented so that they are not left behind once the polymer scaffold degrades after implantation. Imperfections like gels can create an edge on the inner lumen surface that could result in blood flow turbulence. Platelets can also attach to the imperfections, which can lead to thrombosis (clotting of the blood). Additionally, gels within the scaffold tubing wall can impact the mechanical integrity of the scaffold. Gels are very difficult to break down along the extrusion screw, and are also difficult to trap in screen packs. Adding a very dense screen pack before the breaker plate is not a very good solution because it will lead to high operating pressures and additional shear heating. Proper filtering must be fairly fine (40 to 50 microns or finer) to have a noticeable effect, and should have a large filtration area to allow adequate run time before a buildup of gels on the filter results in high operating pressures. For instance, a 12mm extruder using a 50mm breaker plate with sintered metal filters will increase the area of filtration by more than 17 times.

The Importance of a Highly Uniform Tubing Wall Thickness

Typical coronary stent scaffolds are extruded in a size range of 2.5mm (0.098") to 3.5mm (.138") diameter with a wall thickness of 150μm (0.006"). The uniformity of the tubing wall thickness (concentricity) along the length of the scaffold is a critical design parameter.

Fig. 3 – Laser machined scaffold. (Credit: Resonetics)

The extrusion of PLLA tubing exhibits minimal polymer chain alignment. Therefore, extruded PLLA tubes are subsequently heated and expanded radially and axially to orient the polymer chains to increase the mechanical properties of the PLLA tubing along the direction of expansion. Radial expansion increases the radial strength of the tube and therefore the radial strength of the final scaffold. Expanding the PLLA tubing also decreases the wall thickness of the tube significantly. It is critical to the expansion process that the extruded PLLA tubing have a very uniform initial extruded wall thickness to ensure uniformity of the mechanical properties after expansion. Following the expansion process, the wall of extruded PLLA and PLGA tubing is laser machined to create mesh-like scaffold pattern of a stent. (See Figure 3)

The laser machining process is sensitive to variation in the thickness of the tubing wall. Therefore the tubing must have a uniform wall thickness along its length prior to and after expansion for high-quality and repeatable laser processing with accurate strut dimensions.

High variation in wall thickness is undesirable since highly uniform wall thickness throughout the extruded tube provides uniform expansion and high yields during laser cutting, and improves the opportunity for more uniform circumferential vessel wall support. For most medical tubing used in delivery systems, a concentricity level of 85 to 90 percent is acceptable. However, for a drug eluding bioresorbable scaffold, it is highly desirable to extrude the PLLA tubing with a concentricity level of 95 percent or higher, which requires specialized automatic concentricity control techniques that have been proven to work successfully for high-viscosity bioresorbable polymers.


Bioresorbable scaffolds that degrade over time will continue to grow and eventually replace permanent metallic scaffolds. New micro extrusion technologies have been developed that protect the molecular weight of the bioresorbable polymer so that the physical properties of the scaffold are maximized after implantation into the body. Micro extruders for implant applications must also be designed for easy cleaning and for operating under cGMP conditions. Additionally, the development of automatic die centering technologies that enable the extrusion of ultra-uniform wall thickness scaffold tubing will have an impact on device performance and provide for smoother process validations related to secondary expansion and laser processing operations.

This article was written by Stephen D. Maxson, Director, Global Business Development —Medical, American Kuhne, York, PA. For more information, Click Here " target="_blank" rel="noopener noreferrer">