BIOS–Research Doctorate School in BIOmolecular Sciences
Ph.D. in BIOMATERIALS (2012-2014)
Laboratory of Bioactive Polymeric Materials for Biomedical and Environmental Applications (BIOlab) – Department of Chemistry and Industrial Chemistry (University of Pisa)
Development of polymeric vascular connection device for
small diameter vascular by-pass
Phd Candidate Giancarlo Lupi M.D.
Department/Laboratory/Institution:
Laboratory of Bioactive Polymeric Materials for Biomedical & Environmental Applications (BIOlab) Department of Chemistry & Industrial Chemistry
Nuclear Medicine Department
Cardiac, Thoracic and Vascular Department Neurosurgery, Neuroscience Assistential Department
Supervisor: Prof. Federica Chiellini, Prof. Mauro Ferrari Tutor: Eng. Dario Puppi, M.D. Anna Paola Erba,
2 Dedicated to Lorenzo, Vittoria, Niccolò and Daniela
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AIM OF THE WORK ... 7
OUTLINE OF THE WORK ... 9
CHAPTER I ... 10
REVIEW OF SUTURELESS ANASTOMOSIS ... 10
1. INTRODUCTION ... 10
2. HISTORICAL REVIEW OF SUTURELESS ANASTOMOSIS ... 11
3. DEVICE EVALUATIONS ... 19 3.1.PINS ___________________________________________________ 19 3.2.WALL EVERSION ____________________________________________ 19 3.3.WALL SQUEEZING ___________________________________________ 20 3.4.ADHESIVES _______________________________________________ 20 3.5.WELDING _________________________________________________ 21 3.6.TUBES AND STENTS __________________________________________ 22 4. ANASTOMOTIC BIOMECHANICS ... 26
5. EXPERIMENTAL ANALYSIS OF TENSILE FORCE OF INDIVIDUALIZED STENTS FOR MICROVASCULAR ANASTOMOSES ... 28
6. CONCLUSION ... 29
7. REFERENCES ... 30
CHAPTER II ... 37
PREPARATION AND CHARACTERIZATION OF BIODEGRADABLE POLYMERIC STENTS FOR SMALL DIAMETER VASCULAR BY-PASS ... 37
1. ABSTRACT ... 37
2. INTRODUCTION ... 38
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3.1.PREPARATION OF SILICONE REPLICA OF CEREBRAL ARTERIES ____________ 40 3.2.MATERIALS _____________________________________________ 41 3.3.FABRICATION OF PHBHHX AND PCL STENTS ____________________ 41 3.4.THERMAL CHARACTERIZATION _______________________________ 42 3.5.COMPRESSIVE MECHANICAL CHARACTERIZATION ________________ 43 3.6.STATISTICAL ANALYSYS ___________________________________ 44
4. RESULTS AND DISCUSSION ... 45
4.1.THE SILICON REPLICA _______________________________________ 45 4.2.COMPUTER-AIDED WET-SPINNING APPARATUS __________________ 45 4.3.STENTS FABRICATION _____________________________________ 50 4.4.MORPHOLOGICAL ANALYSIS _________________________________ 52 4.5.THERMAL CHARACTERIZATION _______________________________ 56 4.6.MECHANICAL CHARACTERIZATION ____________________________ 57 5. CONCLUSIONS ... 60 REFERENCES ... 61 CHAPTER III ... 64
IN VITRO BIOLOGICAL CHARACTERIZATIONS OF POLYMERIC VASCULAR CONNECTION DEVICE FOR SMALL DIAMETER VASCULAR BY-PASS ... 64
1. ABSTRACT ... 64
2. INTRODUCTION ... 65
3. MATERIALS AND METHODS ... 67 3.1.MATERIALS _____________________________________________ 67 3.2.FABRICATION OF PHBHHX AND PCL STENTS ____________________ 67 3.3.IN VITRO BIOLOGICAL CHARACTERIZATION ___________________ 67 3.4.STERILIZATION AND NAOH TREATMENT ____________________________ 67 3.5.CELL CULTURING AND CELL SEEDING _____________________________ 68 3.6.CELL VIABILITY ___________________________________________ 69
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3.7.CELL MORPHOLOGY __________________________________________ 69 3.8.GRAVIMETRIC ANALYSIS OF THE THROMBUS _________________________ 70 3.9.STATISTICAL ANALYSIS ___________________________________ 70
4. RESULTS AND DISCUSSION ... 71
4.1.STENTS FABRICATION _____________________________________ 71 4.2.IN VITRO PRELIMINARY BIOLOGICAL EVALUATION _____________ 72 4.3.CELL VIABILITY AND PROLIFERATION ____________________________ 72 4.4.STENTS THROMBOGENICITY EVALUATION ____________________________ 76 5. CONCLUSIONS ... 77
REFERENCES: ... 79
CHAPTER IV ... 81
A PRELIMINARY EX-VIVO AND IN-VIVO STUDY OF SUTURELESS MICROVASCULAR ANASTOMOSIS WITH ENDOLUMINAL RE-ABSORBABLE STENT COUPLED WITH THE LASER WELDING TECHNIQUE. ... 81
1. ABSTRACT ... 81
2. INTRODUCTION ... 82
3. MATERIALS AND METHODS ... 83
3.1.EX VIVO SURGICAL CHARACTERIZATION SET UP _______________ 84 3.2.IN VIVO SURGICAL CHARACTERIZATION ______________________ 88 4. RESULTS ... 89 4.1.EX VIVO EXPERIMENT ________________________________________ 89 4.2.IN VIVO EXPERIMENT _____________________________________ 93 5. CONCLUSIONS ... 95 REFERENCES: ... 97 CHAPTER V ... 99
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1. OVERALL CONCLUSIVE REMARKS AND FUTURE PERSPECTIVES. ... 99 2. FUTURE PERSPECTIVE ... 102
AIM OF THE WORK
“There is nothing like looking, if you want to find something. You certainly usually find something, if you look, but it is not always quite the something you were after.”
John Ronald Reuel Tolkien
Despite the huge progress of microsurgery, microvascular anastomoses remain a challenging and time consuming procedure with a steep learning curve for the surgeon.
Since the beginning of bypass surgery, proximal and distal anastomoses have been done with hand-held sutures basically found on the principles of suture technique described by Alexis Carrel in 1902 [1].
The comfort to surgeons in performing a reliable anastomosis with the suture technique and the excellence of its long-term results has led to its adoption as the gold standard. But it is a matter of fact that surgical dexterity is still a determinant factor for anastomosis outcome. The surgical failure is mainly related both to the endothelial damage and to the occlusion time of the recipient arteries. Therefore, surgeons need an alternative way to construct a bypass in order to reduce the technical demand, standardize the quality of the surgical procedure, avoid unnecessary arterial clamp, reduce the individual surgical skill as the determinant factor for anastomosis outcome and possibly, expediting the procedure.
As a consequence, Authors’ efforts are focused in exploring new techniques to shorten the time spent in connecting vessels. We can summarize the previously reported strategies into two groups: sutureless strategies and suture delivery devices. Belong to the first group any device such: memory shape metal alloy, biologic or synthetic glue combined or not with scaffold, laser welding with or without organic solder and endo- exo- luminal re-absorbable stents. The second group consists instead of mechanical devices that can release at once multiple stitches, clips or staples in a predetermined geometry. Hitherto all the reported strategies failed in terms of reproducibility. The consequence is that still today
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some patients affected by complex aneurysms, ischemia or cerebral tumours remain untreatable despite those promising works.
On the basis of this background, the aim of the present project is to investigate a new strategy to be used during the procedure of micro vascular anastomosis (< 3 mm in diameter). In the present PhD activity the set-up a combination of technologies is mandatory to face this challenge. A multidisciplinary strategy, which combines polymeric intra-luminal temporary vascular connection system and the laser welding technique is undertaken. First the vascular connector should hold the vessels end together and then the diode laser should complete a watertight suture of the arteries with few or zero stitches.
The first step of research activity is dedicated to develop and compare different biodegradable stents made of different biopolymer: Poly(ε-caprolactone) (PCL) and Poly[(3-hydroxybutyrate-co-3-hydroxyhexanoate)] (PHBHHx). The micro-structured polymeric stents will be investigated and optimized, together with the processing conditions for the fabrication of the micro-structured polymeric stents. A novel computer-aided wet-spinning apparatus allowing for the fabrication of tubular micro fibrous structures will be designed and assembled by combining a translating extrusion spinneret with a rotating mandrel immersed in a coagulation bath.
Subsequently we will dedicate the activity to stents physical-chemical, in vitro and in vivo characterization.
The physical-chemical characterization will include a thermal (thermogramivetric analysis – TGA, differential scanning calorimetry - DSC) and compressive mechanical (Dynamic Mechanical Analyzer, DMA) investigation.
In vitro biological investigations aimed at assessing the cytocompatibility and hemocompatibility of the prepared stents will be performed using Fibroblast (balb/3T3 clone A31) and HUVEC cell culture and gravimetric analysis of the thrombus respectively.
Preliminary in vivo experiments will be performed to assess the suitability of the produced stents to withstand the surgical manoeuvres needed to insert them into carotid artery.
OUTLINE OF THE WORK
The work performed in the accomplishment of the PhD Thesis focused on the development of a new polymeric vascular re absorbable connection device has been implemented in 4 chapters whose content is summarized by following:
1. Review on Sutureless devices: this chapter includes a literature review of all devices for vascular anastomoses designed and constructed over the years; the research provides the basis from which we started. In particular, we performed a careful analysis of the advantages and benefits described for each class of device, divided according to their prevailing mechanism of action. Some basic biomechanics principles of intravascular tubes and stents are listed.
2. Preparation and characterization of biodegradble polymeric stents for small diameter vascular by pass: this chapter describes the preparation and chemical and physical characterization of the wet-spun biodegradable polymeric stents. The processing parameters for the production of small calibre tubular polymeric devices were investigated.
3. In vitro Biological Characterizations of Polymeric Vascular Connection Device for Small Diameter Vascular By-pass: This chapter is focused on the behaviour of HUVEC and Fibroblast cells when seeded on PCL and PHBHHx stents. Several in-vitro assays are taken into consideration to evaluate cell attachment and proliferation.
4. Preliminary ex-vivo and in-vivo study of sutureless microvascular anastomosis with endoluminal re-absorbable stent coupled with the laser welding technique: preliminary “ex-vivo” and “in-vivo” evaluation of the stents prototypes are presented to assess their suitability to withstand the surgical manoeuvres needed to insert them into carotid artery.
CHAPTER I
REVIEW OF SUTURELESS ANASTOMOSIS
1. INTRODUCTION
« A few observation and much reasoning lead to error; many observations and a little reasoning to truth. » Alexis Carrel
This chapter aims to assess the relevant Literature for micro-vascular sutureless techniques. In fact, we review the anastomotic devices developed, focusing on the key issues of sutureless anastomosis and demonstrating that the ‘new anastomotic technologies’ are based on concepts expressed in the 19th century. In particular an overview of the major sutureless strategies adopted over the years is given to establish the context of the contributions.
In the first part we build a timeline of the procedures developed and reported over the years, and we compare them to the Gold Standard of the hand-sewn technique. We especially focused on results in terms of: pitfalls, advantages and reproducibility of the technique described by each Author.
In the second section we classify the methods previously described according to their prevalent mechanism of action. This chapter is therefore a prerequisite to our research and technical and technological innovations.
The chapter ends with a review of the existing surgical strategies that are nowadays described in a clinical environment.
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2. HISTORICAL REVIEW OF SUTURELESS ANASTOMOSIS
Alexis Carrel was a surgeon and biologist who has contributed to fundamental advances in the techniques of suturing blood vessels and the research on transplantation of tissues and organs, essential for daring surgical operations of our time. He received the Nobel Prize in 1912 for the outstanding achievements in his work on vascular anastomoses, and even today the technique described by Carrel is considered the gold standard [1]. During his Nobel lecture he affirmed: Many surgeons had previously to myself performed vascular anastomosis, but the results were far from satisfactory. He was referring to vascular non-suture techniques that had been described since the end of the 19th century. His words witnessed the huge effort that the Authors have made to achieve more efficient vascular anastomoses in all epochs. Those works lead to non-suture techniques and new devices for proximal and distal anastomoses that represent our base. In fact we should read achievements and limits between the strategies proposed in the frame of biomedical innovations of our time.
Even if we consider the end of the 19Th century as the starting point of the
research in the field of sutureless anastomoses, we would remind that it date back to the Neolithic period. In fact, according to Biomaterials Science, at that epoch in South Africa and India, biting ants were used to repair wounds.
If the first attempts at revascularization start 300 years after Christ when trying Cosmas and Damian re-implanted a leg repairing vessels with needles ivory. In 1774 Le Conte replaced a femoral artery with a quill pen using a sutureless technique [2].
Is a matter of fact that since the 1890s, experimental studies of connecting blood vessels raised and Authors investigated both non-suture and suture techniques. Abbe [3] introduced first intraluminal glass prosthesis in 1894 (Figure 1) and later: ivory cuffs [3], paraffined silver tubes [4] and the - of an ox [5]. These endo and exo-vessels devices were far from good but constitute the foundation for subsequent developments.
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Figure 1 – Abbe glass prostheses: it acts as a connector between the severed vessels. The prosthesis is filled by water before put it in place. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
In 1900, Payr [6] presented an absorbable extraluminal magnesium ring (Figure 2). The proximal vessel was passed and everted through the ring to complete a circumferential ligature. The distal end was dilated for insertion of the rigid cuff with its everted vessel. The anastomosis was completed by another circumferential ligature, thereby achieving intima-to-intima apposition.
Figure 2 - Payr's magnesium rings: the severed vessels passes through the ring and it is then everted; the others vessel’s end is dilated wrapped around the first one. The first concept ring needed two circumferential ligatures to secure the anastomosis. The second proposed ring had pins that avoided the ligatures. Goetz replicated the same design in 1961. Goetz’s ring was made by tantalum instead of magnesium. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
In 1902, Carrel described the use of dissolvable intraluminal stents composed of caramel candy cylinders for vascular anastomosis1, but this technique has never been applied clinically.
In 1904, Payr [7] presented two interlocking magnesium rings. One of the two rings had pins that allowed to keep the proximal vessel everted. The distal vessel
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was dilated and passed outside the first one to envelop the ring. The second ring sealed the construct obtained an intima – intima contatc surface. Unfortunately Payr’s results were far from those obtained with Carrel’s technique.
Blakemore [8] introduced vein graft-lined rigid vitallium tubes to bridge arterial defects (Figure 3). In an extensive review of vascular trauma during that Second World War,
Figure 3 – Blackmore vitallium tube: different methods of usage are shown. The same technique is replicated for both the severed vessels. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
De Bakey and Simeone reported a slight increase in the number of amputations required following Blakemore’ s technique, although it was not statistically significant [9].
Between 1945 and 1950, a group of Russian engineers and physicians developed a mechanical apparatus for vascular anastomosis. Inverted U-shaped tantalum clips with pointed ends were used to perform end-to-end anastomoses in vessels ranging from 1,3 to 20 mm in diameter.
The first report on its clinical application came in 1956 from Androsov [10], who did not use the device for anastomosis but for the repair of various vessel injuries and traumatic aneurysms (Figure 4).
The same device was used and then modified by Androsov and Inokuchi in order to obtain end-to-end and end-to-side anasomoses.
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Figure 4 – The staple device proposed by Androsov. The device was able to deliver multiple staples at once in a pre-ordinated geometry. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
Swenson and Gross [11] in 1947, used absorbable fibrin tubes in 27 end-to-end jugular or caval vein anastomoses in dogs. After 6–7 weeks the fibrin tubes had dissolved, leaving patent anastomoses in 26 specimens. The tubes Weiss and Lam used for end-to-end femoral artery anastomoses in dogs were made of tantalum [12]. They achieved a patency rate of 86 per cent in vein grafts between two short tantalum cuffs, but attempts with longer tantalum tubes were less successful. In 1953, Bikfalvi and Dubecz [13] applied silver clips by means of a vessel encircling mechanism based on the principle of the von Petz clamp [14].
In 1961 Goetz completed the first internal mammary– coronary artery anastomosis in dogs by a mechanical device. The device resembles the Payr’s rings even if was made by tantalum instead of magnesium. Goetz long-term results were superior to direct suture methods on beating heart [15].
In 1962, Nakayama [16] repaired vessels range between 1.5–4.0 mm in diameter by the usage of a ring-pin device with instrumentation (Figure 5).
Figure 5 – Nakayama’s ring. The severed vessels were everted through the rings that own 6 pins and 6 holes each ones. The pins were intended to enter the reciprocal hole into the paired ring. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
In 1965, Ota and colleagues [17] used a soluble gelatin stent to perform vascular anastomoses. Although the stent dissolved, it temporarily impeded blood flow by swelling.
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Fibrin glue was introduced in 1977 by Matras et al. [18] as a material for the creation of vascular anastomoses. They described their results with fibrin sealed end-to-end carotid artery anastomoses in rats. This was a safe and effective method, provided that two 10/0 stay sutures remained permanently. The disadvantages for the usage of the fibrin glue, allergic reactions and anaphylaxis, pushed the Authors to abandon this path [19, 20].
The Food and Drug Administration (FDA) in the USA cancelled licences for manufacture of fibrinogen in 1978, as the heat treatment, at that time required to inactivate hepatitis B virus, would damage the fibrinogen product [21].
Gottlob and Blumel [22]. in 1968, used alkyl-cyanoacrylates to secure bushings that were used for experimental vascular anastomosis in vessels ranging from 1,0 to 5,0 mm in diameter. They reported satisfactory short-term patency rates, but already stated in the same paper that they had changed the type of cyanoacrylate because of histotoxicity. Other group tested the 2-octyl cyanoacrylate with good results [23].
In the 1980s, Kirsch and associates [26, 27] developed a non-penetrating method of vascular anastomosis, in which small titanium clips were applied to everted vessel edges in an interrupted fashion.
In 1984 an absorbable anastomotic coupler device for vessel ranged between 1 to 2 mm was presented (Figure 6). It had a patency rate of 96% at 12 months in animals but its diffusion was limited by a mean vessel sacrifice of 4 mm in length for vessel’s eversion over the cuffs [28].
Figure 6 - Daniel's device: anastomotic coupler device for vessel ranged between 1 to 2 mm. It consist of three elements: two symmetric anvil the hold the everted vessel’s ends and a central connector that block the others element together. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
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Based on the Nakayama apparatus, Berggren introduced the Unilink system (3M Company, St. Paul, Minnesota, USA) in 1986 [29].
The heat-shrink tubing was introduced in 1991 to ensure the connection between the edges of the severed vessel (Figure 7) [30].
Figure 7 – Mattox’s technique: the heat-shrink tubing was introduced in 1991 to ensure the connection between the edges of the severed vessel, everted onto extravascular anvil. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
In 1994, Moskovitz et al. [31] described a fibrin glue-based microvascular anastomosis applied over a soluble stent made of monoglycerides, diglycerides and triglycerides. The stents were used in vessels ranging from 0·3 to 0·8 mm in diameter. Although the technique was faster than suturing, late patency rates were reduced owing to aneurysm formation.
In 1999, a Japanese group of investigators published their experimental results with a pin-ring coupler with absorbable rings that consisted of L-lactid acid and glycolic acid [32, 33, 34]. Although the rings dissolved through hydrolysis, the stainless steel pins persisted. Furthermore, the chronic inflammatory reaction,
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especially in the adventitia, was more intense than that in sutured anastomoses [32].
In 2000, a new intravascular stent made of a metal alloy of nickel and titanium was designed for sutureless end-to-side anastomoses between coronary arteries and grafts. The T-shaped device (the Graft Connector, Jomed International AB, Helsingborg, Sweden), expands after its insertion into the receiving vessel [35]. In the same year Gundry et Al. [36]. described a new approach for biologic glued anastomoses, based on an intraluminal scaffold that preserve the 3D anatomic relationship of the vessels (Figure 8). In short two catheters with inflatable balloons were inserted into the vessels to be anastomosed. The expanded balloons acted as intraluminal scaffold allowing the surgeon to strew the anastomosis of glue. According to the Authors, this approach could prevent the vessels from collapse e displace when the glue solidifies. The technique has never reached the pre clinical phase.
Figure 8 - Gundry’s technique: endovascular balloon act as a scaffold and the anastomoses is completed by the use of biological glue. (P.Tozzi. Sutureless Anastomoses, secrets for success. Springer, 27 mag 2007)
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The presented devices present limitations that have prevented their full acceptance into the mainstream of vascular surgery. The main limitations observed could be summarized into few key points:
- complex surgical instrumentation and procedures;
- non elastic endo – exo vascular devices conflict with the biological elasticity of the arteries;
- mismatch between size or shape of the device and arteries;
- arterial tissue consumption resulting in high tensile stress of anastomosis; - retention strength not sufficient with no stitches
To overcome these limitations new devices or combination of them should be developed to obtain adequate surgical reliability.
In more recent years several proximal and distal anastomotic devices have been commercialized so we need to summarizes the most interesting systems available in both experimental and clinical settings. Most of the devices are discussed in the next chapter.
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3. DEVICE EVALUATIONS
Any vascular reconstruction achieved without the hand-sewn technique may be considered as a mechanical sutures. Therefore, more than 80 patents can be listed in the area of the sutureless vascular anastomoses, even if most of them are still far from a clinical use.
Starting from the devices previously liested, we will attempt to group them according to their prevailing mechanism of action. Pros and cons reported are presented.
The prevailing mechanisms of action of those devices are listed:
1) pins,
2) wall eversion 3) wall squeezing 4) adhesive 5) welding 6) tube and stents
Those solutions have been proposed also in combination.
3.1. PINS
Pins can link the connector to the vessel [5] passing through the arterial wall. Since the pins are directed inside-out and their intimal damage is comparable to that of a needle of size 5-0. The tension force is limited just by the wall’s resistance that can sometimes tear. Any intrinsic damage of the wall contraindicated their usage. We need to stress that in case of rupture of the anastomosis, the edges of the vessel must be trimmed with substantial loss of vascular tissue and higher resulting tensile force.
3.2. WALL EVERSION
The vessels stumps could be everted on an anvil and it ensures the juxtaposition of the intima – intima surface as dictated by Carrel. Tensional strength is excellent,
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but there are several potential limitations such the hardness of the wall that could lead to an intimal lesions mostly on the edge. The first generation devices almost ever resulted into the necrosis of the everted vessel at their junction. Subsequent device’s improvement showed better results for this specific issue on histological studies obtained from animal implants [37].
3.3. WALL SQUEEZING
The device should squeeze the vessel wall between his inner and outer surface, avoiding the wall’s trauma and preserving the vasa vasorum by covering a very small intimate surface. The thickness of the wall should be between 2 and 4 mm depending on the device’s design, otherwise the squeezing force is insufficient resulting into an anastomosis failure.
The tensional force is lower than that observed with sutures but in case of excessive tearing there is no need to trimmer the vessel edges.
Again the rules dictated by Carrel are respectd by the wall squeezing [1].
3.4. ADHESIVES
The adhesives used clinically may be categorized into two groups: a) fibrin glues
b) cyanoacrylate glues.
Fibrin glue consists generally of two components. The first one contains fibrinogen, factor XIII and plasma proteins whereas the second one is composed of thrombin, aprotinin and calcium chloride. Firstly reported on 1977 [18], the fibrin glue’s usage is a safe and effective method provided that few stitches remain permanently. It is time saving and is easy to apply [38] but it present some drawbacks. In addition to the narrowing of the lumen at the anastomotic site the cons include allergic reactions and anaphylaxis [19]. Therefore the Food and Drug
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Administration (FDA) in the USA cancelled licenses for manufacture of fibrinogen in 1978 [21].
The synthetic cyanoacrylates (methyl-, ethyl- and butyl- cyanoacrylates) came to the Authors attention to prevent the cons reported for the fibrin’s derivate but further reports on the tissue toxicity of cyanoacrylates when implanted around blood vessels limited their usage. Authors noted an early marked foreign body granulomatous response with later thinning of the vessel wall, splitting of the elastic lamina and calcification of the media [39, 40]. The development of false aneurysms at the anastomotic join has also been reported [42]. The mechanisms liable for these results might be the high heat of that comes from the polymerization and the excessive amount of glue applied. Recent reports suggest that 2-octyl-cyanoacrylate may be less toxic, but the FDA has approved it just for topical use in wound closure [43]. A proper adjustment of the vessel stumps is mandatory to prevent the glue from entering into the vessel lumen, so stitches are anyway required.
In summary, fibrin adhesives or cyanoacrylates for vascular anastomosis have little advantage and several drawbacks, especially allergic and toxic reactions. A further problem is that a few sutures continue to be required to prevent leakage or aneurysm formation in the longer term.
3.5. WELDING
Since 1979 [44] several types of lasers have been explored for vascular anastomoses: argon [45], carbon dioxide [46] and excimer with diode probe [47]. Even if the strength of the anastomosis performed with lasers has been improved through the use of solders [48], the thermal damage and long hemodynamic stress with consequent anastomotic failure is still the main challenge [49]. Moreover, the potential spreading of the solder (albumin e.g.) into the arterial lumen may impair the vascular patency. Authors therefore investigated combined solutions in which the laser should be coupled with a different endovascular device. For example an endovascular catheter would prevent the spreading of the solder through the anastomosis as shown on coronary goat [50].
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Another technical advancement comes from the recent introduction of a thermal detection systems directed to assess the vessel wall’s temperature, to check for the optimal laser’s parameters in case of different types of seal.
In short, laser welding both for end-to-end and for end-to-side vascular anastomosis has mainly been used experimentally and no clinical application has yet been established.
Disadvantages include:
- poor anastomotic strength, - expensive hardware
- specific surgeon’s learning curve to improve operative knowledge of instrumentation
- The technique is not standardized and results are therefore operator’s related.
3.6. TUBES AND STENTS
The use of a tube or stent to perform a vascular anastomosis date back to 1894 [2]. The first dissolvable intraluminal stent report , made of caramel candy, was in 1902 [1]. Later Authors tested other biomaterials, among these ones the fibrin tubes [11], tantalum [12], soluble gelatin stent [17] need to be mentioned. The subsequent approach was based on water-soluble tube: saccharide mixture [51], mono-, di- and tri- glycerides [31]. Those tubes were coupled with fibrin glue to complete the vascular joint but results were far from good. In fact, tubes made by tantalum, soluble gelatin and fibrin resulted into a swelling with the consequence of hypoperfusion. The other mentioned biomaterial resulted into an insufficient patency rate because of aneurysm formation.
In more recent times the contact surface’s area between the blood and the tube became an important issue to reduce the thrombogenicity. It has been observed that the more this area is wide, the more is the risk of thrombogenicity.
Other reported causes of failure are [52]: poor vessel walls apposition, poor suture tension, uneven suture spacing, over manipulation of the stent with subsequent intimal damage [53].
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Since the 1978, when the first catheter balloon was patented, intra-vascular stents became the wider field of research. In fact, the use of stents is not just about open surgery but also and especially for endovascular procedures. The massive experience acquired about the benefits and cons using the endovascular stents is a base for our research devoted to the new surgical purpose of vascular anastomosis. More recently, the development of bioabsorbable stents has become a hot topic in the medical device industry [54], so our starting point could be the huge knowledge that industry and clinicians state about those endovascular devices. Models of endovascular bioabsorbable stents that are currently being developed are made of either polymers or corrodible metal alloys.
There are few polymeric bioabsorbable stents into clinical trials: The Igaki-Tamai coronary stent [55]
the bioabsorbable everolimus-eluting coronary stent [57]
the stent made of tyrosine poly(desaminotyrosyl-tyrosine ethyl ester): REVA Medical stent [60]
sirolimus stent from Bioabsorbable Therapeutics [69]
The Igaki-Tamai coronary stent and the bioabsorbable everolimus-eluting coronary stent both use Poly-L-lactic acid (PLLA).
The Igaki-Tamai stent is the first bioabsorbable stent to be implanted in humans, has a thickness of 0.17 mm and is balloon-expandable [55]. Full degradation took 18-24 months. Furthermore, at about 36 months, lumen size increased [56].
The everolimus-eluting bioabsorbable PLLA stent has: a thickness of 150 µm and a polymer coating that release everolimus, a drug that decrease blood supply to reproducing cells. Full absorption of the stent is completed over 18 months [57]. It showed a higher restenosis rates [58] compared with the not reabsorbable ones, but a remarkable finding is that, between 6 months and 2 years an enlargement in lumen size is documented [59]. This result suggests that after stent absorption, the vessel could recover its health again.
The REVA Endovascular Study of a Bioresorbable Coronary Stent (RESORB), with a thickness of 150 µm, is coated with paclitaxel [60]. The stent is
balloon-24
expandable and is set into place by sliding and locking parts rather than deforming the material. Between 4 and 6 months after implantation, there was higher-than-anticipated occurrence of repeated percutaneous coronary intervention, driven mainly by reduced stent diameter [59].
The polymeric stent developed by Bioabsorbable Therapeutics is coated with sirolimus, an immune system suppressing drug. The stent has a thickness of 200 µm and is balloon-expandable. Both the base polymer and coating polymer of the stent are made up of bonds between salicylic acid molecules. When the bonds are hydrolyzed the release salicylic acid is released and it could theoretically prevent the vessel from restenosis. Full absorption is expected within 6 to 12 months. A significant thickening of the inner lining of the artery has been documented
Metal alloy bioabsorbable stents perform similarly to permanent metallic stents. Magnesium stents have potential advantages over polymeric stents in terms of higher radial strength.
Metal alloy bioabsorbable stents perform similarly to permanent metallic stents. So far, two bioabsorbable metal alloys have been proposed for this application:
• iron
• magnesium
Neither of these stents is coated with drugs.
Magnesium stents have potential advantages over polymeric stents in terms of higher radial strength. The first magnesium stent implanted in humans has a thickness of 165 µm and is balloon expandable [59]. In trial, absorption of a magnesium stent in humans was rapid and mechanical support lasted days or weeks, which is too short to prevent restenosis [57, 58]. During the first four months, major adverse cardiac events were recorded in 24% of the patients and after one year up to 45% of the patients had additional percutaneous re-treatment. The magnesium stent can be safely degraded within 4 months, but the high restenosis rate raises concerns [57].
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The experimental iron stent has a thickness of 100–120 µm and is balloon-inflatable. The researchers implanted stents made of 41 mg of pure iron, an amount equivalent to the monthly oral intake of iron for a human, into the descending aortas of New Zealand white rabbits. During the 6 to 18 months of follow up, there was no reported thrombosis or any other significant inflammatory injury response. However, the animals experienced destruction of the internal elastic membrane of arteries and products from the degradation of the stent accumulated, resulting in significant alteration of the artery wall [61].
The main purpose of the described stent was to solve arterial stenosis in case of atherosclerotic plaque or inflammatory wall diseases. Even if they represent an huge experience about the arterial wall reactions, we have few experience in case of readsorbable stent temporary technique in micro vascular anastomosis. Up to date the Authors are aware only a stent made of nickel and titanium developed for sutureless end-to-side anastomoses [35].
In summary, Polymeric biodegradable stents have demonstrated several limitations and long-term effects of polymer full absorption products are unclear The polymer that both the Igaki-Tamai and BVS stents use, PLLA, holds 1,000 mmHg of crush pressure and maintains radial strength for approximately one month. Compared with metallic stents, this radial strength is lower and may result in early recoil post-implantation [57]. Meanwhile, metal alloys stents do not seem biocompatible enough to use in practice.
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4. ANASTOMOTIC BIOMECHANICS
Elastic properties of vessel wall are altered by vascular anastomoses: the more the anastomosis is compliant, the less is the probability that it could develop a stenosis due to myointimal hyperplasia [62].
There are several studies that underline the importance of anastomotic compliance in the vascular reconstruction outcome [63]. Unfortunately, there is no experimental study on sutureless anastomotic compliance.
The balance between the stiffness of the system and the time during which it is exerted must be considered to reduce the risk of chronic injury of the vessel wall. In fact the conflict between a rigid foreign body and a dynamic dilating artery could result into severe vessel wall atrophy [64].
The effect of graft calibre (i.e., donor to recipient diameter ratio) is another parameter evaluated to assess the hemodynamics of anastomosis. It is observed that donor to recipient diameter ratios larger than 5:3 have better performance than smaller ones [66]. Besides, smaller grafts typically present an increased risk of early graft failure due to thrombosis [67].
The effect of competitive flow (i.e., flow through a bypassed native artery) on the graft patency has been extensively investigated, but still is somewhat controversial.
The majority of the investigations on biological factors are in vivo studies along with complementing in vitro investigations. On the other hand, computational simulations of blood flow along with numerous in vitro and in vivo investigations constitute the studies of the biomechanical factors and hemodynamic parameters, and have provided strong evidence on the influence of these factors on initiation and onset of intima hyperplasia.
Also the choice of the correct anastomotic angle between donor and recipient arteries is mandatory to reduce the generation of vortex.
Authors reported that the intimal hyperplasia of the arterial wall of the anastomoses is strictly correlated to the graft angle, and graft patency rates vary according to flow velocity in proximity to the wall. This branch angle represents a relatively common in vivo situation and is therefore a convenient starting point.
27
Computer fluodynamic simulations [33] and haemodynamic studies [34] suggest that the best angle for a side-to-end or end-to-side reconstruction is between 30 and 45.
In conclusion, there are a few aspects to be considered in the design of an optimal anastomosys, but since a comprehensive discussion of the biomechanics of the anastomosis is not among the purposes of this thesis we therefore refer to the principal recommendations reported in the literature [72]:
- Compatibility of the graft with the arterial pressure and the supplied blood flow rate, to ensure a physiologic range of intramural stresses and
hemodynamic forces in the graft itself.
- Arterial compliance of the graft, to avoid compliance mismatch with the host artery at the anastomotic junction, to prevent escalation of intramural stresses in the artery and the graft, which can result in IH formation, especially on the suture-line.
- Hemodynamic performance driven design of anastomotic configuration of the distal anastomosis, to regulate the hemodynamic parameters and wall shear stress indices, in order to avoid triggering of the pathogenic factors of IH and thrombosis (e.g., platelet activation, long near-wall residence time, etc.).
- Minimal vascular injury, to minimize proliferation of SMCs as a wound healing response. Technological advances may further develop the suggested alternatives to sutures (e.g., biological glues, laser generated solders, etc.) to a practicable level for routine clinical use. Not only can such products minimize vascular injury, but also they can eliminate the para-anastomotic hyper-compliant zone and the associated elevating intramural stresses which are caused by the stiff sutures.
- Patient-specific designs, to tailor the design considerations to each particular patient’s vascular characteristics.
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5. EXPERIMENTAL ANALYSIS OF TENSILE FORCE OF INDIVIDUALIZED STENTS FOR MICROVASCULAR ANASTOMOSES
Overall, more than 80 patents prove the huge interest of the Authors to simplify and shorten the procedure in case of vascular anastomosis. The main reported pitfalls include leakages from the anastomoses, the early occlusion of the vessel, and the necessity for additional sutures for implant fixation.
Since Julio Palmaz patented the concept of balloon expandable stents in 1985 [73] few studies demonstrated the safety and feasibility of stent in case of open artery anastomosis, many of which are based on permanent metallic platforms (stainless steel or cobalt–chromium) and therefore not absorbable with potential long term adverse reaction [68, 69].
One of the major challenges for stent-based anastomoses is their mechanical properties. Normally, no intravasal tensile force is applied to the stent. Use of the stent as a connection device requires tensile forces, which is a major challenge in the provision of a tight, leak-proof, and reliable anastomosis.
Therefore authors investigated the influence of length, dilation and differences in fabrication of the newly developed balloon-expandable stent on the tensile force of stented anastomoses [2]. The presented preliminary tests established that anastomoses sutured by a single stitched technique with eight stitches can hold approximately two-thirds the tensile force of a non-sutured vessel. In that work the Authors referred a satisfactory mean force of 180 N for the stented anastomosis if compared with the 8 stitches one of 96 N (long 24 mm stents, dilated approximately 1 mm larger than vessel diameter). Starting from that basis they suggest to better suite the mechanical, technical, and material characteristics of individualized stents and catheters in case of stented microvascular anastomosis in free tissue transfer.
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6. CONCLUSION
Sutureless vascular anastomosis is a procedure gaining interest for realizing rapid and easy surgeries. Most of the anastomotic devices presented may be employed for large-calibre vessels, and there are still no devices for the small-calibre ones (< 2 mm diameter) [70].
Several anastomotic experimental devices based on different techniques have been proposed but unfortunately they never reach the mainstream of surgery, probably because of technical and engineering limitations of those years.
Old concepts should be so reassessed in the light of new technologies, biomaterial evolution and better understanding of vascular anastomosis. In this respect, the present research activity attempts to develop an intravascular stent that could act as a vessel connector. This strategy is undertaken by a multisciplinary approach comprising biomaterials science and engineering and clinical practice.
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CHAPTER II
PREPARATION
AND
CHARACTERIZATION
OF
BIODEGRADABLE POLYMERIC STENTS FOR SMALL
DIAMETER VASCULAR BY-PASS
1. ABSTRACT
This research activity was aimed at the preparation and characterization of biodegradable polymeric stents as endovascular connection device for small diameter vascular by pass. For this purpose, a novel computer-aided wet-spinning apparatus allowing the design and additive manufacturing of polymeric stents was developed. The processing parameters for the production of small calibre tubular polymeric devices were investigated and tailored also on detailed solid replica of the cerebral arteries (circle of Willis).
By employing the optimized production parameters, poly(ε-caprolactone) (PCL) and poly(3-hydroxybutyrate-co-3-hydroxyexanoate) (PHBHHx) stents with tailored size and porosity were developed. Prototypal samples with different wall thickness and porosity were characterized for their morphological, thermal and mechanical properties. PCL stents showed significantly larger wall thickness and pore size in comparison to PHBHHx stents. Thermal analysis results suggested that the employed technique did not cause remarkable changes in chemical-physical properties of the polymers. PHBHHx stents demonstrated great radial elasticity and PCL stents showed higher mechanical strength, comparable with other polymeric stents produced by means of other techniques
38
Keywords: biodegradable stents, wet-spinning, poly(ε-caprolactone), poly(3-hydroxybutyrate-co-3-hydroxyexanoate)
2. INTRODUCTION
Bioabsorbable stents have emerged in the past years as one of the most promising approach in obstructive cardiovascular disease treatment due to their potential in providing mechanical support while it is needed and then disappearing from the vessel, allowing its natural healing and avoiding the risks associated with having a permanent metallic cage [1, 2]. A number of bioabsorbable stents made of polymeric or metallic materials have been investigated over the last years. Together with magnesium alloys, aliphatic polyesters are by far the most employed biodegradable polymers for stents development due to their relatively long degradation rate and suitable mechanical properties. In particular, p oly(L-lactic acid) (PLLA), poly(glycolic acid), poly(D,L-oly(L-lactic-co-glycolic acid) (PLGA), and poly(ε-caprolactone) (PCL) have been investigated for bioabsorbable stents development [3, 4].
The Igaki-Tamai stent (Kyoto Medical Planning, Kyoto, Japan) was the first resorbable stent implanted in humans [5]. It consists of an extruded high molecular weight PLLA filament with a zig-zag helical coil design. The stent received the CE marking in 2007 to treat peripheral artery disease (PAD), and recent clinical outcomes showed its long term safety as coronary scaffold that could disappear within 3 years [6]. The ABSORB vascular scaffold (Abbott Vascular, CA, USA) is a drug-eluting stent based on PLLA loaded with everolimus, a drug with antiproliferative action [7]. The ABSORB received the CE Mark in 2011, and the device is now available in more than 60 countries worldwide. Other biodegradable stents currently under clinical evaluation are the REVA Medical stent made of tyrosine polycarbonate (REVA Medical Inc., CA, USA), the IDEAL stent (Bioabsorbable Therapeutics Inc, Menlo Park, CA) made of a polyanhydride and salicylic acid, and the novolimus-loaded DESOlve stent made of PLLA (Elixir, CA, USA) [8, 9].
Wet-spinning is a polymeric fibre fabrication process based on non-solvent induced phase inversion, which involves the extrusion of a polymeric solution directly into a coagulation bath [10]. The polymeric solution precipitates because
39
of the solvent/non-solvent counter-diffusion that lowers polymer solubility, leading to the formation of a continuous polymer fibre. Solvent/non-solvent demixing causes the formation of a polymer-rich phase and a polymer-lean phase that usually results in a sponge-like fibre morphology. In the past years, 3D scaffolds composed of an assembly of wet-spun fibres made from different biodegradable polymers such as PCL, PLLA, chitosan and starch, were investigated for tissue engineering applications [11-13]. However, the methods developed to assemble the wet-spun fibres into a 3D scaffold don’t allow achieving good structural reproducibility and production efficiency. A computer-aided wet-spinning (CAWS) technique based on the controlled deposition of a wet-spun polymeric solution for the layered manufacturing of 3D objects was developed to overcome these limitations. Recent studies reported on the development of 3D scaffolds made of PCL [14], PCL with a star molecular structure [15], or PHBHHx [16] by CAWS achieving good control over internal architecture and external shape.
The aim of this study was the development of biodegradable stents for small diameter vascular by-pass with special emphasis for cerebral arteries. The cerebral arteries, unlike splanchnic or cardiac arteries, have short straight sections. These traits are the only ones suitable to perform a bypass. As a consequence any stent has to be ideated also according also to this peculiar profile. As a consequence we first developed a detailed solid replica of the cerebral arteries (circle of Willis), starting from a multi-detector CT scan dataset (MDCT). Segmentation and 3D model generation were performed using a semiautomatic tool developed in the EndoCAS centre [17]. The quantitative geometrical and topological data allowed improving the comprehension of anatomy and the dimensions of cerebral arteries. The stent connector was so tailored according also to this information by a novel CAWS apparatus. The CAWS apparaturs was designed and assembled for the production of 3D microstructured polymeric constructs with a tubular geometry. The processing parameters for the production of those small caliber stents made of PCL or PHBHHx were so optimized according and the developed prototypal samples were subsequently characterized for their morphological, thermal and mechanical properties.
40
3. MATERIALS AND METHODS
3.1. PREPARATION OF SILICONE REPLICA OF CEREBRAL ARTERIES
For this purpose, a semi- automatic tool, the EndoCAS Segmentation Pipeline [17] integrated in the open source software ITK-SNAP 1.5 (www.itksnap.org) [18] was used. The whole segmentation procedure is based on the neighbourhood connected region-growing algorithm that, appropriately parameterized for the specific anatomy and combined with the optimal segmentation sequence proposed generates good-quality 3D images coupled with facility of use. A pourable platinum curing silicone rubber was employed to fabricate the synthetic arteries. The silicone (Dragon Skin® series by Smooth on) (DSFxPro) was used in combination with a liquid rubber for casting with a cold cure by poly-condensation. The DSFxPro is suitable for the manufacture of small and detailed objects. The arterial tree was fabricated using the 3D printer (Dimension Elite 3D printer), starting from the previous segmentation. The DSFxPro was then smeared on the assembled mould. After complete silicone curing, the solid replica of the arteries was removed so that the arterial lumen was Preserved. The silicon replica presents the same arterial size and shape of the arterial tree, including the lumen size (Fig. 1). The scaffolds geometry could be so tailored and tested before the “ex vivo” and “in vivo” phase. A pourable silicone rubbers and an agarose gel were employed to fabricate the synthetic arteries.
Figure 1 Image of the cerebral artery silicon replica. It resembles the intra cranial carotid and Sylvian axis.
41
3.2. MATERIALS
Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx, 12 mol% HHx, Mw = 300000 g/mol) was kindly supplied by Tsinghua University (China). Poly(ε-caprolactone) (PCL, CAPA 6800, Mw = 80000 g/mol) was supplied by Perstorp Caprolactones Ltd (Warrington, Cheshire, UK). All the solvents were purchased from Sigma-Aldrich (Italy) and used as received.
3.3. FABRICATION OF PHBHHX AND PCL STENTS
PHBHHx was dissolved in chloroform under stirring for 2 h at 30 °C to obtain a homogeneous solution at the desired concentration. PCL was dissolved in acetone under stirring for 2 h at 35 °C to obtain a homogeneous solution at the desired concentration. For the preparation of PHBHHx and PCL stents, the prepared polymeric solution was placed into a glass syringe connected to a blunt tip stainless steel needle through a plastic tube. By using a programmable syringe pump (KDS100, KD Scientific, MA, USA), the solution was injected at a controlled feeding rate directly onto a rotating mandrel immersed in an ethanol bath. The X-Y movement of the needle and the mandrel rotation around its axis were controlled by an in-house made computer-controlled system allowing for the predefined winding of the coagulating wet-spun polymeric fibre around the mandrel (Fig. 2). The prepared stents were left under a fume hood at room temperature for 24 h and subsequently vacuum-dried to remove any residual solvent.
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Figure 2. Scheme of computer-aided wet-spinning (CAWS) apparatus for the fabrication of tubular microfibre structures.
3.4. THERMAL CHARACTERIZATION
Thermal properties of raw and processed polymers were obtained by Thermogravimetric Analysis (TGA) and Differential Scanning Calorimetry (DSC). TGA was carried out using a TGA Q500 instrument (TA Instruments, Italy) in the temperature range 30 – 400 °C for PHBHHx and 30 – 600 ºC for PCL, at a heating rate of 10 ºC/min and under a nitrogen flow of 60 mL/min. Degradation temperature (Tdeg) was evaluated as the temperature corresponding to
a percentage weight loss of 2%. DSC analysis was performed at a heating rate of 10 ºC/min, cooling rate of 20 °C/min and under a nitrogen flow rate of 80 mL/min, using a DSC-822E instrument (Mettler Toledo, Italy). PHBHHx samples were analyzed in the range –40° to 180 ºC, while PCL samples in the range -100 to 100 ºC. Melting temperature (Tm) and crystallinity degree (C%) were evaluated
43
3.5. COMPRESSIVE MECHANICAL CHARACTERIZATION
Compressive mechanical properties of both PHBHHx and PCL stents were evaluated using a DMTA V machine (Rheometric Scientific, Germany). Five specimens for each kind of stents were tested at a strain rate of 1 %/s [19] until a maximum strain of 80 %. Tubular samples were placed between the two testing parallel plates and tested in either axial or radial compression (Fig. 3 and 4).
Figure 3. Axial compression of a stent CAWS at the beginning and end of the test; the yellow arrow indicates the direction of the compression.
Figure 4. Radial compression of a stent by CAWS at the beginning and end of the test; the yellow arrow indicates the direction of the compression.
The reaction force was defined as the measured force, whilst the strain was evaluated as the ratio between the stent length/diameter variation and its initial length/diameter for axial/radial compression, respectively. In order to evaluate the capability of the developed stents to recover their initial length/diameter when subjected to a load, a strain rate of 1 %/s was applied up to a given strain and soon after load release the length/diameter of the sample (LE) was measured. Initial length/diameter recovery (LR) was calculated as the percentage ratio between LE
