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shape, and they found that collapse pressure increased firstly within 12 weeks and then deceased. From the
material point of view, the initial increase of compressive properties might be caused by the recrystallization
of biodegradable polymer PLA/PLLA due to the water absorption and temperature increase (from room
temperature 20 °C to human body temperature 37 °C). It was reported that crystallization can significantly
improve the mechanical strength and stiffness of PLLA. It should be pointed out that there is a lack of efforts
in assessing the discontinuity of the scaffold during the late stage of degradation, due to the difficulties in
imaging the scaffold. However, such studies can be accomplished by using the micro computed tomography
(CT) technology to investigate the internal damage associated with both biodegradation and mechanical
deformation wherever feasible. Also the microstructures of all specimens can also be investigated using SEM
both before and after the tests, by fracturing the specimens (e.g., in liquid nitrogen) and coating the fracture
surfaces with a thin layer of gold.

In addition, clinical trials have also been carried out extensively and long-term follow up indicated the safety
and efficacy of polymeric stents in treatment of coronary artery disease. Follow-up studies, based on invasive
imaging methods such as CT, angiography and intravascular ultrasound (IVUS), were normally carried out to
assess the stented vessel after the implantation of bioresorbable scaffolds. As reported in Ormiston et al. , the
[27]
bioresorbable vascular scaffold (BVS) absorb showed an overall 16.8% reduction of lumen area at 6-month
follow-up of 30 patients implanted with the scaffolds. Serruys et al. evaluated the outcomes of Absorb for
[28]
treating coronary artery stenosis in 45 patients. The mean lumen area was found to decrease by 3.1% (6.60 to
6.37 mm ) according to IVUS analysis at 6-month follow up whereas it increased to 6.85 mm at 2-year follow
2
2
up. Recent 5-year follow-up results after Absorb implantation concluded that the mean lumen area tended
to increase from 6 months to 1 year and 5 years . Progressive lumen gain was also observed after Absorb
[29]
implantation in porcine coronary arteries , for which the lumen area kept stable up to 6 months and then
[30]
showed a progressive increase from one to 3.5 years. These studies confirmed the lumen gain after scaffold
implantation over the degradation times, especially from 6 months onwards.
However, studies are still limited and challenging due to the microscale geometry of stent and complex
environment of human artery. To complement experimental work, finite element simulations were
consequently used to study the stent performance during deployment, which is reviewed in the next section.

COMPUTATIONAL WORK
Stent deployment
Finite element (FE) modelling is of particular help in evaluating stent performance. However, most FE
modelling of stent implantation, including recent ones, were dealing with metallic stents such as the effects
of design and material on stent expansion, recoiling and dogboning [31-33] . FE modelling of polymeric BRSs
is very limited, especially how they compare with metallic stents in terms of mechanical performance.
There are only a few papers in literature. For instance, Pauck and Reddy simulated the mechanical
[34]
performance of PLLA stents with different designs and varying polymer stiffness. Material stiffness and
geometrical design affects the radial strength of polymeric stent significantly. Debusschere et al. studied
[35]
the free-expansion behaviour of bioresorbable Absorb stent by considering the viscoplasticity of the
material. They simulated stent expansion by increasing inflation pressure linearly or in a stepwise, and
linear inflation method was found to induce higher stress concentration in the U-bend struts. Inflation
rate affected stent expansion mainly during inflation process, while not much difference was shown after
balloon deflation. Wang et al. studied the mechanical deformation of a PLLA biodegradable coronary
[36]
stent using experimental and computational methods. They modelled the crimping of stent to an outer
diameter of 1.41 mm and subsequent expansion to different inner diameters. Stent deformation profiles
were in good agreement with experimental ones during both crimping and expansion processes, and large
deformation occurred at sharp curvature of struts in all cases. The radial recoil ratio showed a decrease