During the last two decades, a great amount of researches have been focused on biodegradable metals. Technologies from alloy design to melting, manufacturing and processing, from micro-tube to stent laser processing and drug eluting coating have been improved and optimized continuously. Biodegradable metallic stent has evolved from a concept to a real product and generated three branches of material system. A large amount of animal tests and clinical tests have been carried out to investigate biodegradable magnesium stents. Results of clinical study have indicated that the magnesium stent is feasible, with favourable safety and performance outcomes. More importantly, Biotronik won CE Mark for Magmaris bioresorbable stent in 2016. Researches of biodegradable iron stents are still in the stage of animal tests. The nitrided iron stent possesses excellent mechanical properties. Results showed a good long-term biocompatibility of nitrided iron stent in rabbit and porcine model. Biodegradable zinc stent has only been introduced in recent years. Only a few in vivo studies have been reported with zinc wires implanted in rats. Results showed a good degradation behavior and biocompatibility of zinc wires. In this paper, the current research status of biodegradable metallic stents is reviewed, and the future research and development in mechanical property optimization, drug eluting and intelligence is proposed.
Fund: Supported by National Key Research and Development Program of China (No.2016YFC1102402) and National Natural Science Foundation of China (No.51431002)
Fig.1 Schematic diagram of degradation behavior and the change of mechanical integrity of biodegradable metallic stents during the vascular healing process[1]
Fig.2 SEM images (upper panels) and EDX mapping of AMS-3.0 degradation products (lower panels, unrelated to upper panels) (At 28 d, non-degraded magnesium-alloy particles had been surrounded by magnesium, calcium, and oxygen, probably in the form of MgCO3 or Mg(OH)2 or both, which could not be differentiated without further analysis. At 90 d, magnesium-alloy area was reduced, while oxidated (yellow) areas had become partly replaced at their outer margins by a calcium-phosphorous-oxygen compound (bluish area). Raman and infrared spectroscopy combined with X-ray diffraction analysis clarified that this was calcium phosphate with amorphous structure. At 180 d, no remaining metallic particles were noted)[23]
Criterion
Constraint
Biodegradation
Mechanical integrity for 3~6 months; Full absorption in 12~24 months
Biocompatibility
Non-toxic and non-inflammatory; No allergenic potential; No harmful
strength : elastic modulus ratio>0.16; Elongation to failure>15%~18%;
Elastic recoil on expansion<4%
Microstructure
Homogeneous and approximately isotopic
Small grain size
<30 μm
Corrosion rate
Penetration rate<0.02 mma-1
Table 1 Summary of material criteria and constraints for a biodegradable stent[9]
Test
Stent system
Experiment
Biocompatibility
Degradation
Ref.
model
time
Animal
AE21
Domestic
No thromboembolic events, 40% loss of
89 d
[17]
test
pigs, coronary
lumen diameter corresponding to
artery
neointimal formation, 25% re-enlargement
caused by vascular remodeling resulting
from the loss of mechanical integrity
between days 35 and 56
AZ91
Dogs,
Lumen was clear and no elastic recoil and
7 d
[18]
coronary or
thrombosis, moderate intimal hyperplasia
femoral artery
at 14 d
AZ31B
Rabbits,
Lumen area was significantly greater, the
120 d
[19]
P(LA-TMC)+
abdominal
neointimal area was significantly smaller
sirolimus
aorta
and endothelialization was delayed at 30 d
in coated group
WE43
Minipigs,
Inhibitory effect on the smooth muscle
98 d
[20]
coronary
cells, rapid endothelialization, thin layer of
artery
neointima covering the stent after 6 d,
degradation caused inflammation and
intimal hyperplasia
AMS
Pigs,
No signs of ongoing inflammation,
2 months
[21]
coronary
smallest lumen area at 3 months because
artery
of negative vascular remodeling
AMS
Pigs,
Safe and with less neointimal formation
-
[22]
coronary
compared with stainless stent, lumen area
artery
did not change
AMS-3.0
Pigs,
Equivalent to TAXUS Liberte regrading
180 d
[23]
PLGA+
coronary
late luminal loss, intimal area, fibrin
paclitaxel
artery
score and endothelialization. Inflammation
score was high at 28 d but disappeared at
later time
Clinical
AMS
Preterm baby,
No relevant inflammatory reaction to the
5 months
[24]
study
pulmonary
stent material, minimal alteration of the
artery
vessel wall and an increase of the arterial
diameter after stenting
AMS
Newborn,
15 d after implantation, blood velocity
-
[25]
aortic arch
increased significantly, blood perfusion
recovered, lumen diameter increased
from 1.5~1.8 mm to 2~2.8 mm
AMS
20 patients
The clinical patency rate was 89.5% after 3
-
[26]
months, no blood toxicity was found
PROGRESS-
63 patients
No myocardial infarction, subacute or late
4 months
[27]
AMS
thrombosis, or death. Angiography at 4
months showed an increased diameter
stenosis of 48.4. Overall target lesion
revascularization rate was 45% after 1 a.
Neointimal growth and negative
remodeling were the main mechanisms
of restenosis
Biosolve-I
46 patients
Target lesion failure was 7% at 12 months.
-
[28]
DREAMS,
A significant reduction of lumen area at 6
PLGA+
months and 12 months follow-up. No cardiac
paclitaxel
death or scaffold thrombosis
Biosolve-II
123 patients
A preservation of the scaffold area with a
12 months
[8,29]
DREAMS 2G,
low mean neointmal area. Target lesion
PLLA+
failure was 4%. No definite or probable
sirolimus
scaffold thrombosis was observed. QCA
parameters remained stable from 6 months
to 12 months. Target lesion failure was
3.4% at 12 months
Table 2 In vivo tests of biodegradable Mg based stents[17-29]
Fig.2 SEM images (upper panels) and EDX mapping of AMS-3.0 degradation products (lower panels, unrelated to upper panels) (At 28 d, non-degraded magnesium-alloy particles had been surrounded by magnesium, calcium, and oxygen, probably in the form of MgCO3 or Mg(OH)2 or both, which could not be differentiated without further analysis. At 90 d, magnesium-alloy area was reduced, while oxidated (yellow) areas had become partly replaced at their outer margins by a calcium-phosphorous-oxygen compound (bluish area). Raman and infrared spectroscopy combined with X-ray diffraction analysis clarified that this was calcium phosphate with amorphous structure. At 180 d, no remaining metallic particles were noted)[23]
Stent system
Animal model
Biocompatibility
Exp. period
Ref.
Iron
Rabbits,
No thromboembolic complications, no
6~18 months
[35]
descending
adverse events. No significant neointimal
aorta
proliferation, no pronounced inflammatory
response and no systemic toxicity
Iron
Pigs,
No signs of iron overload or iron-related organ
360 d
[36]
descending aorta
toxicity, no local or systemic toxicity
Iron
Pigs, coronary
At 28 d, no stent particle embolization or thrombosis
28 d
[37]
artery
and no excess inflammation, or fibrin deposition
Iron
Rats, artery
Substantial corrosion at 22 d, a voluminous
9 months
[38]
lumen or wall
corrosion product retained within the vessel
wall at 9 months. Implant in artery lumen
experienced minimal corrosion
Nitrided iron
Minipigs, iliac
Endothelialization after 1 month. Slightly lumen loss
12 months
[39]
artery
at 12 months. No thrombosis or local tissue necrosis
Nitrided iron Zn+
Rabbits,
Complete endothelialization after 3 months,
13 months
[40]
PDLLA+
abdominal
slight inflammation during implantation,
sirolimus
aorta
no necrosis and systemic toxicity
Iron, nitrided iron
Rabbits,
Endothelialization after 7 d. Slight
53 months
[7]
abdominal
inflammation during implantation. No
aorta,pigs,
necrosis and systemic toxicity. Corrosion
coronary artery
products can be cleaned by macrophages
Table 3 In vivo tests of biodegradable Fe based stents[7,35-40]
Fig.3 SEM, EDS (a) and micro-CT (b) images of iron based scaffold (IBS) after 6 months implantation in rabbit abdominal aorta[34]
Implant
Animal model
Biocompatibility
Experimental
Ref.
period
Zinc
Rat, abdominal
Retained about 70% of its original cross
6 months
[46]
wire
aorta wall
sectional area after 4 months, after which
degradation was observed to increase
rapidly. Corrosion products consisted
of ZnO, ZnCO3 and trace of Ca/P
Zinc
Rat, abdominal
A complete endothelial layer at 2.5 months
6 months
[47]
wire
aorta wall
and stable appearance at 6.5 months.
Smooth muscle cells remained stable at
6.5 months, no pronounced chronic
inflammation
ZnAl
Rat, abdominal
No acute and chronic inflammatory were
6 months
[48]
wire
aorta wall
presented, no necrosis. Cross-section was
reduction 40%~50% at 6 months
ZnAl
Rat, abdominal
Inflammatory cells were able to penetrate
6 months
[49]
wire
aorta wall
the corrosion layer of ZnAl implant. A
delayed entrance of inflammatory cells
into corrosion layer of pure Zn was observed
ZnLi
Rat, abdominal
Degradation rates were 0.008 and 0.045 mm/a
12 months
[50]
wire
aorta wall
at 2 and 12 months, respectively. No neointimal
hyperplasia. Inflammation and neointima
thickness was slightly higher for ZnLi than Zn
Zinc
Rat, abdominal
Intense fibrous encapsulation of the wire, steady
20 months
[51]
wire
aorta wall
corrosion without local toxicity for up to 20
months. Chronic inflammation at 5~8 months
but subsided between 10~20 months
Zinc
Rabbit,
No severe inflammation, platelet aggregation,
12 months
[52]
stent
abdominal
thrombosis formation or obvious intimal
aorta
hyperplasia was observed
Table 4 In vivo tests of biodegradable Zn wires[46-52]
Fig.4 Zn-H2O (a) and Zn-C-H2O (b) Pourbaix diagrams for physiological concentrations at 310 K (The dotted lines show physiological pH of 7.4. The physiological potential for tissue fluid is indicated by circles. E—potential)[51]
Fig.5 Schematic diagrams showing the evolution of degradation mechanism of zinc stent associated with the conversion of degradation microenvironments during healing process(a, b) formation of zinc phosphate under the dynamic flow condition in blood fluid(c, d) conversion of zinc phosphate to ZnO and calcium phosphate under the diffusion condition in neointimal. SEM images corresponding to the related schematic diagrams are consisted of representative surface morphologies and cross-sections. The timeline depicts the healing process including inflammation, granulation and remodeling phases and selected time points. The models explain the formation of corrosion products and their dependence on local pH, distance from stent surface and implantation time in blood fluid (e) and neointimal (f). The red lines represent the assumed pH variation on the sample surface[52]
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