Influence of design parameters on the mechanical behavior and porosity of braided fibrous stents

    Influence of Design Parameters on the Mechanical Behavior and Porosity of Braided Fibrous Stents Rita Rebelo, N´ıvea Vila, Raul Fangueiro, Sandra Carvalho, Sohel Rana PII: DOI: Reference:

S0264-1275(15)30121-0 doi: 10.1016/j.matdes.2015.07.051 JMADE 279

To appear in: Received date: Revised date: Accepted date:

20 March 2015 9 July 2015 10 July 2015

Please cite this article as: Rita Rebelo, N´ıvea Vila, Raul Fangueiro, Sandra Carvalho, Sohel Rana, Influence of Design Parameters on the Mechanical Behavior and Porosity of Braided Fibrous Stents, (2015), doi: 10.1016/j.matdes.2015.07.051

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ACCEPTED MANUSCRIPT Influence of Design Parameters on the Mechanical Behavior and Porosity of Braided Fibrous Stents

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Rita Rebeloa,b,1, Nívea Vilaa, Raul Fangueiroa,SandraCarvalhob,cand Sohel Ranaa Fibrous Materials Research Group - Centre for Textile Science and Technology, University b

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of Minho, 4800-058 Guimarães, Portugal.

GRF-CFUM - Physics Department, University of Minho, 4800-058 Guimarães, Portugal. c

SEG-CEMUC Mechanical Engineering Department, University of Coimbra, 3030-788

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Coimbra, Portugal.

Abstract

In spite of all innovations in stent design, commonly used metallic stents present several

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problems such as corrosion, infection and restenosis, leading to health complications or even death of patients. In this context, the present paper reports a systematic investigation on

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designing and development of 100% fibre based stents, which can eliminate or minimize the problems with existing metallic stents. For this purpose, braided stents were produced by varying different material, structural and process parameters such as mono-filament type and

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diameter, braiding angle and mandrel diameter. The influence of these design parameters on mechanical behaviour as well as stent’s porosity was thoroughly investigated, and suitable

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parameters were selected for developing a stent with mechanical characteristics and porosity matching with the commercial stents. According to the experimental results, the best

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performance was achieved with a polyester stent designed with 0.27 mm monofilament diameter, braiding angle of 35° and mandrel diameter of 6 mm, providing similar properties to commercial Nitinol stents.

Keywords: Fibrous stents, Braiding, Design parameters, Mechanical behavior, Porosity.

1

[email protected] Universidade do Minho, Campus de Azurém, 4800-058 Guimarães, Portugal.

ACCEPTED MANUSCRIPT 1.

Introduction

Cardiovascular diseases cause 40% of annual deaths in the European Union and lead to an estimated cost of 196 billion Euros [1]. Atherosclerosis of coronary arteries is one of the most

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common cardiovascular diseases [2]. This type of disease can be treated either by drugs like anticoagulants or antiplatelet agents or by surgical procedures. Cardiovascular stents are used

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in the surgical procedure called “angioplasty” and are tubular and perforated structures. They are inserted into the blood vessels to keep the vessels open and to reestablish the normal blood flow. For application as implantable medical devices, stents should possess a few important characteristics such as mechanical strength, radiopacity, longitudinal flexibility,

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high radial expansion, corrosion resistance, biocompatibility and easy handling [3].

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Stents can be made of different materials but metallic stents, particularly stainless steel, are the most common commercial stents due to their good mechanical characteristics. However, metallic stents possess some demerits such as low radiopacity, corrosion, restenosis and

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bleeding complications. Restenosis occurs in approximately 30% of the patients after the surgical procedure and is associated with a high mortality rate as well as high cost of health

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care [3]. One of the main problems caused by restenosis is the damage of arteries by stent’s metallic rods, resulting in an immune and inflammatory response by the human body [4]. In

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order to minimize the problems associated with the metallic stents, new materials like fibres are now being explored. As the surface of fibrous materials can be easily functionalized to enhance their compatibility with human tissues, they present high potential for application in

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stents [5]. Furthermore, monofilament fibres exhibit excellent mechanical properties and perform very well when subjected to compression, tensile and bending forces, and therefore, can fulfill the stent’s requirements related to mechanical performance [6].

Fibrous stents may be produced using woven, knitting and braiding technologies. Braiding technique is the most suitable technique for stent development as the produced structures exhibit good dimensional stability, mechanical strength, high shape recovery and flexibility [4]. Moreover, braiding technique is a simple and versatile process, providing the structures high compliance level, torsional stability and wear resistance [4]. Braided structures have been well studied for application in medical implantable devices like sutures, artificial ligaments, etc [7, 8]. A few studies have demonstrated the possibility of using braided structures for producing self-expanding stents. For example, Zhao et al. [9]

fabricated

braided stents using shape memory Nitinol wires. The developed braided stents, however,

ACCEPTED MANUSCRIPT showed lower radial strength as compared to the welded stents, owing to relative sliding of Nitinol wires within the braided structures and therefore, found less effective to restore the patency of the stenotic artery. Further studies on Nitinol braided stents revealed that due to

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the slippage between the wires and superelastic behaviour of Nitinol, they showed a hysteresis between loading and unloading cycles under the compressive loading [10].

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Braiding angle was observed to be the main design parameter deciding the mechanical performance of Nitinol braided stents. Braided stents have also been fabricated using shape memory polymers such as polyurethane and their mechanical performance has been characterized and simulated using finite element analysis [11] . It was observed that the

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design parameters like fibre diameter and braiding angle can be carefully selected to match the performance of polyurethane stents to that of metallic stents. However, shape memory

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polyurethane shows viscous phenomena such as stress relaxation and creep, which should be minimized. Also, due to much inferior mechanical properties of this polymer to Nitinol, it should be used as thicker fibres, reducing blood flow through the vessels. The strong

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influence of number of filaments and stent diameter on the mechanical performance of stents was also noticed in case of braided stents produced with polyester fibres [11]. Braided stents

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fabricated using polypropylene fibres showed better mechanical performance as compared to knitted stents, indicating the advantage of braiding technology for stent development [6].

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Possibility of producing bioabsorbable braided stents has been explored using PLLA fibres [7] and the produced stents showed similar radial pressure stiffness to metallic stents and maintained, at least, half of their radial pressure stiffness for more than 22 weeks. These

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initial research studies on braided stents investigated separately the influence of some of the design parameters on selected properties of stents such as radial compression. However, effective designing of braided stents should consider the influence of all important design factors related to material (such as fibre type and diameter), structure (such as braiding angle) and process (mandrel diameter) on the various required characteristics such as porosity and cover factor, radial and axial compression, bending behaviour, etc. As no existing study investigated all these factors together, a systematic investigation is highly essential to study and understand all these factors and properties, in order to produce suitable fibrous stents for replacing commercial metallic stents. To address this, the present research investigates the influence of different design parameters such as monofilament type and diameter, braiding angle, and mandrel diameter on the porosity and various mechanical properties (longitudinal and radial compression and bending properties) of braided stents and suitable parameters

ACCEPTED MANUSCRIPT have been selected to design fibrous stents with characteristics matching with the commercial Nitinol stents.

Materials

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2.

2.1. Monofilaments

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Polyester (PES), polyamide (PA) and polypropylene (PP) monofilaments were used to produce braided stents. Linear density, mechanical and frictional properties of these monofilaments in three different diameters were evaluated. For the determination of single fibre braking force and elongation at break, ISO2062 standard was used. Linear density was

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measured following NP2060 standard. The results of raw material characterization are

2.2.Fabrication of braided stents

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provided in Table 1.

In order to produce the braided stents, a braiding machine with 16 yarn bobbins was used.

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Braided structures were formed due to interlacing of monofilament yarns in the diagonal directions forming integrated structures. The monofilaments were interlaced around a

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mandrel, which defined the final stent diameter.

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Table 1-Properties of raw materials

Diameter (mm) Linear density

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PES

PA

PP

0.27

0.35

0.55

0.30

0.40

0.50

0.26

0.30

0.37

82.79

133.38

341.45

73.49

126.00

214.49

67.09

71.10

101.30

39.07

58.30

142.66

40.52

72.20

107.80

28.86

43.65

47.74

(2.84)*

(2.17)

(1.35)

(1.64)

(2.16)

(4.13)

(1.16)

(4.60)

(0.68)

5.79

1.10

4.65

2.03

1.75

1.72

2.57

5.51

2.88

(2.13)

(2.70)

(1.96)

(5.44)

(7.86)

(8.74)

(2.38)

(1.86)

(6.40)

(Tex) Tensile strength (N) Initial Modulus( N/Tex)

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0.44

0.42

0.56

0.57

0.50

0.43

0.61

0.47

(N/Tex)

(2.84)

(2.17)

(1.35)

(1.64)

(2.16)

(4.13)

(1.16)

(4.63)

(0.68)

Elongation

16.15

33.44

19.07

30.48

30.01

30.59

27.36

15.24

21.12

(%)

(9.78)

(4.69)

(6.78)

(7.76)

(5.55)

(13.73)

(21.04)

(8.61)

(5.37)

Frictional

0.40-

0.40-

coefficient

0.45

0.45

0.3-0.35

0.3-0.4

0.35-

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Density

0.25-0.3

0.45

1.14

* The values within the brackets are C.V.(%)

0.15-0.20 0.25-0.3

0.9

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1.38

(g/cm3)

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Tenacity

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All stents were produced using 16 filaments. Stents with different characteristics were produced by varying the monofilament diameter, mandrel diameter (3.2, 5 and 6 mm) and

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braiding angle (20º, 30º and 35º). Braiding angle is an important parameter which influences directly the mechanical performance. This parameter depends on the number of braiding yarns and the take-up speed. The braiding angle () represents the angle between the

α

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monofilaments and the longitudinal axis of braided structure, as shown in Figure 1 [12].

Figure 1- Braided structure showing braiding angle. In the present study, braiding angle was calculated using Eq. 1 [13]:

Braiding angle,

Where, α= braiding angle (rad); w= angular speed of bobbins (rad/s);

(1)

0.120.15

ACCEPTED MANUSCRIPT R= mandrel radius (cm); v= take-up speed (cm/s).

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Angular velocity and angle made by horn gears can be calculated using the following

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equations:

Angular velocity,

Where,

(2)

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θ= angle made by a pair of adjacent horn gears at the machine center (rad);

(3) (4)

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t= time taken by the horn gear to make one revolution (s).

Where,

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Nh= number of horn gears;

Therefore,

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Wh= angular velocity of the horn gears around their own centers (rad/s).

(5) (6)

For easier understanding of the tests and results, sample coding was done according to material type, monofilament diameter, braiding angle and diameter of mandrel. Thus, the sample code PES270A20D32 corresponds to a polyester sample with monofilament diameter of 0.27mm, braiding angle of 20° and mandrel diameter of 3.2mm.

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Methods

3.1. Determination of Cover Factor and Porosity The amount of fibrous materials deposited on the mandrel surface during the braiding process

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depends on the braiding process parameters and can be estimated using cover factor. Cover factor is defined as the percentage of mandrel surface covered by monofilaments. This is a

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good indicator of braid’s structural uniformity [13]. Porosity is another important parameter for stents and should range between 70% and 80% [10]. Porosity and cover factor were

Where, Wy=monofilament width (mm); R= mandrel radius (mm);

(8)

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α= braiding angle (rad).

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Nc= number of bobbins;

(7)

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calculated using Eq.7 and Eq.8.

3.2.Measurement of Radial Compression

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It is highly essential to characterize the radial compression of stents, as stents with insufficient radial force can lead to revision surgeries. When a radial pressure is applied, stents undergo a structural deformation leading to an expansion in the longitudinal direction and shrinkage in the transversal direction (Figure 2). This deformation results in change of stent diameter and braiding angle [14].

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Figure 2- Structural changes in a stent when subjected to radial pressure.

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Due to radial compression over time, stents decrease in diameter and increase in length. Therefore, the relationship between the length (Lf) of a single filament per turn and stent

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diameter (D) becomes:

(10)

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Where:

(9)

D= new average stent diameter (mm);

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Da= average stent diameter (mm); α0= Initial braiding angle (rad);

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α= new braiding angle (rad).

The extension in the total length of stent (∆ ) can be obtained by equation 11.

(11)

For the radial compression test, stents were held between two flat plates in a universal testing machine and 15 samples with 8 cm in length were compressed up to 25% of their original diameter, as shown in Figure 3. The average force of compression was then calculated.

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Figure 3– Setup of radial compression test

3.3.Determination of Longitudinal Compression

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Stents should have good axial compression resistance; otherwise they will be easily compressed at their ends due to a slight trauma while exchanging different catheters. Developed stents were compressed to 15% of their original lengths and the required force for

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this compression was measured [15]. For longitudinal compression tests, 15 samples of 8 cm length were compressed to 1.2 cm in a universal testing machine (Figure 4) and the average

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force was calculated.

Figure 4-Setup of longitudinal compression test

3.4. Determination of Bending Properties Bending behavior plays a vital role in deciding the stent performance. After stents are implanted in to the human body, they should maintain at least, 75% of its original diameter, in order to allow blood flow and avoid stent collapse. The diameter of a bent stent was measured at the midpoint of the bent section of the stent (shown in Figure 5 by arrow) according to the method reported in Ref. [16]. For this purpose, 3 samples with 8 cm length

ACCEPTED MANUSCRIPT were bent until their ends were separated by 2 cm (Figure 5). The ratio of the stent diameter in the bent condition and the initial diameter (expressed in percentage) was obtained and this parameter represents the percentage of original diameter which remained unchanged during

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the bending test. Unchanged diameter can be calculated using Eq. 12.

Where:

(12)

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Figure 5- Measurement of unchanged bending diameter

Db= stent diameter in bent condition (mm);

4.

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Di= initial stent diameter (mm);

Results and discussion

4.1. Cover Factor and Porosity 4.1.1. Influence of fibre type and diameter Monofilament diameter strongly influences the cover factor and therefore, the stent porosity. As per equation 7, for the same production parameters, samples with lower monofilament diameter should show a lower cover factor and, consequently, higher porosity. This is attributed to the fact that the filaments with lower diameter (i.e. lower width) can cover less area on the mandrel surface for a particular mandrel diameter and braiding angle. Consequently, samples with higher monofilament diameter should present high cover factor and lower porosity. The influence of monofilament diameter and type on stent’s porosity is shown in Figure 6. It can be noticed that the experimental results are in good agreement with

ACCEPTED MANUSCRIPT Eq. 7. For all materials, an increase in monofilament diameter reduced the porosity due to

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better cover factors.

Figure 6-Influence of fibre type and diameter on porosity [mandrel diameter: 6mm, braiding

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angle: 35⁰].

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Also, the material type influenced the porosity of produced stents mainly depending on their diameters. Overall, due to lower monofilament diameter polypropylene stents presented

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lower cover factor and higher porosity (reaching maximum porosity of 81%) as compared to

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polyamide and polyester stents with maximum porosity of 78% and 76%, respectively.

4.1.2. Influenceof mandrel diameter The influence of mandrel diameter on cover factor of produced stents is shown in Figure 7. It is clear that the cover factor decreased with the increase in mandrel diameter. This is in agreement with the trend which can be predicted using Eq. 7. If all other parameters maintained the same, an increase in mandrel diameter deposits less number of yarn coils on the mandrel and thereby reduces the area of mandrel covered by yarns. Therefore, an increase in mandrel diameter decreases its cover factor and consequently, increases the porosity.

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Figure 7-Influence of mandrel diameter on cover factor of PA stents (a), PES stents (b) and

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PP stents (c) [braiding angle: 35⁰].

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4.1.3. Influence of Braiding Angle

Figure 8 presents the influence of braiding angle on cover factor of produced stents. An

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increase in the braiding angle led to an increase in the cover factor. This trend is also in agreement with Eq. 7. As shown in Figure 1, braiding angle is the angle between the yarn and

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the longitudinal axis of braids and therefore, in increase in the braiding angle leads to deposition of braiding yarns more towards the transverse axis. As a result, if a certain length and volume of mandrel is considered, higher braiding angle resulted in deposition of higher number of yarn coils and therefore, increased the cover factor. Consequently, porosity of the braided stents decreased with the increase in braiding angle.

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Figure 8 - Influence of braiding angle on cover factor of PA stents (a), PES stents (b) and PP

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stents (c) [mandrel diameter:3.2mm].

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4.2. Radial Compression

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Theoretically, the radial compression force of stents can be expressed as:

(13)

Where, P= radial pressure (MPa); F= radial force (N); nc= number of bobbins; D= average stent diameter (mm); L= length of stent (mm); α= braiding angle (rad).

Therefore, theoretically, radial compression force is dependent on the applied force, number of filaments, stent diameter and length as well as braiding angle.

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Figure 9 - Influence of fibre type and monofilament diameter on radial compression force for

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stents produced with 20° braiding angle and 3.2 mm mandrel diameter

4.2.1. Influence of fibre type and diameter The influence of fibre type and diameter on radial compression force is shown in Figure 9.

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The force for radial compression is highly dependent on the monofilament type and their diameter. As the diameter of different monofilaments was not same, it was quite difficult to

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understand the independent effect of monofilament type on radial compression. But overall, polyester stents required higher force than polyamide stents and the lowest resistance to

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radial compression was shown by the polypropylene stents. Also, the force for radial compression strongly increased with the monofilament diameter. This was attributed to several reasons: (a) higher monofilament diameter led to a higher stent wall thickness, resulting in higher radial compression resistance, (b) higher monofilament diameter also improved the cover factor, thereby improving the structural integrity and compactness of stents and this resulted in higher compression resistance, (c) radial compression led to bending and buckling of the monofilaments and therefore, higher bending or buckling resistance at higher diameters also increased the force for radial compression. The radial compression force for polyester samples varied between 16.19N and 1.06N; for polyamide varied from 15.55N to 0.31N and for polypropylene varied from 3.66N to 0.31N. In case of PP stents, no significant difference in radial compression force was observed in three diameters, as the diameter difference was small.

ACCEPTED MANUSCRIPT 4.2.2 Influence of mandrel diameter The influence of mandrel diameter on radial compression force is presented in Figure 10. It can be observed that, except PES550 stent, an increase in mandrel diameter resulted in a

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decrease in the force required for radial compression. As the stents were compressed between two flat parallel plates during radial compression, bending and buckling of the monofilaments

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occurred, besides some structural rearrangements which increased the length of stents. An increase in mandrel diameter also increased the diameter of stents and the length of the monofilaments under the buckling force. Since the buckling force is inversely proportional to the length of monofilaments, an increase in mandrel diameter reduced the compression

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resistance of the stents. However, the specimen PES550 did not follow the trend as shown by the other specimens. This exception was perhaps due to the lower frictional coefficient of

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these monofilaments, which could cause monofilament sliding during the compression test

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and therefore, behaved differently from other samples.

Figure 10-Influence of mandrel diameter on radial force of PA stents (a), PES stents (b) and PP stents (c) [braiding angle:35⁰].

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Figure 11-Influence of braiding angle on radial compression force of PA stents (a), PES

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stents (b) and PP stents (c) [mandrel diameter:3.2mm].

4.2.3. Influence of braiding angle

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The influence of braiding angle on radial compression force is presented in Figure 11. It can be noticed that for most of the stents an increase in the braiding angle increased the radial

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compression force. This trend is in agreement with Eq.13. The primary reason behind this was the deposition of yarns more towards the transverse axis of the stents at higher braiding angles, leading to more resistance towards transverse forces acting during radial compression. Moreover, higher braiding angle also improved the cover factor and compactness of the stent structure resulting in higher resistance towards radial forces. However, too high braiding angle may also reduce the interlocking of the filaments (as they are arranged more parallel to each other) and increase the chance of slippage during compression. As a result, a decrease in the radial force (i.e. variation from regular trend) was also observed for some samples at high braiding angles and with samples presenting low frictional co-efficient such as PES550, PA500 and PP300.

ACCEPTED MANUSCRIPT 4.2.4. Diameter recovery and extension in length As mentioned in Section 3.2, redial compression leads to increase in the length of stents and this can be calculated using Eq. 11. After the compression test and relaxation, the stents

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showed some change in the diameter and length, indicating unrecoverable deformation. Table 2 lists the diameter recovery and longitudinal extension for selected PA and PES samples

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produced with high monofilament diameter and lower mandrel diameter. It can be noticed that after compression test and relaxation, braiding angles (α) changed to some extent from the original ones (α0), indicating structural rearrangement. Some increase in stent’s length can also be noticed. Although, stent diameter could not recover completely back to the original

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values, the diameter recovery was more than 80% (Table 2), which is comparable with the

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value obtained with previously reported polypropylene stents [7].

α0 (⁰)

PA500A20D32

35

PA500A30D32

Da (mm)

D (mm)

Diameter recovery (%)

∆h (mm)

28.48

4.2

3.49

83.10

1.38

27.95

4.3

3.82

88.84

0.68

α(⁰)

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Sample

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Table 2- Diameter recovery and longitudinal extension of stents after radial compression.

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31.02 32.64

29.44

4.2

3.83

91.19

0.7

PES550A20D32

34.29

29.2

4.3

3.72

86.51

1.12

PES550A30D32

30.35

28.94

4.3

4.12

95.81

0.33

PES550A35D32

31.84

25.16

4.3

3.47

80.70

1.42

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PA500A35D32

4.3. Longitudinal Compression The axial force on tubular stents can be expressed by the following equation [14].

(14) Where: ;

ACCEPTED MANUSCRIPT Where, nc= number of bobbins; E= Young’s modulus (Pa);

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G= Shear modulus (Pa); I= Moment of inertia (kg*m2);

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Ip= Polar moment of inertia (m4) α0= Initial braiding angle (rad); α= new braiding angle (rad).

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Da= average stent diameter (mm).

It can be noticed from Eq.14 that, theoretically, the axial force depends on several

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parameters such as number of yarns, modulus of rigidity and elasticity of braided tubes, moment of inertia of the fibers, braiding angle and stent diameter. However, this relationship is quite complex to understand the independent effect of each parameter on

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axial force.

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4.3.1. Influence of fibre type and diameter Figure 12 presents the longitudinal compression force for samples with braiding angle

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of 20° and mandrel diameter of 3.2 mm. It can be noticed that PES stents showed higher axial compression force as compared to PA stents, which exhibited significantly higher axial compression force than PP stents. The higher axial compression resistance of PES

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stents was attributed to the higher modulus of polyester filaments as compared to polyamide filaments. The main reasons for much lower axial compression force of PP stents were the lower diameter of PP monofilaments as well as their lower modulus. Stents containing higher diameter monofilaments exhibited higher longitudinal compression force due to their higher bending rigidity, as during longitudinal compression stents are also subjected to bending and buckling forces. Similar to the case of radial compression, PP stents developed with different monofilament diameters presented very small difference in longitudinal compression behaviour, mainly due to small difference in their diameters.

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Figure 12- Influence of fibre type and diameter on longitudinal compression force

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[braiding angle: 35⁰, mandrel diameter: 3.2mm].

4.3.2. Influenceof mandrel diameter

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The relationship between longitudinal compression force and mandrel diameter can be observed in Figure 13. Similar to radial compression, an increase in the mandrel

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diameter led to a decrease in the longitudinal compression force. Higher mandrel diameter resulted in lower cover and inferior structural integrity of the stents, resulting in lower resistance towards axial compression force.

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Figure 13- Variation of longitudinal force with diameter of mandrel of PA stents (a),

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PES stents (b) and PP stents (c) with braiding angle of 35⁰.

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4.3.3. Influence of braiding angle

Figure 14 presents the influence of braiding angle on longitudinal compression force.

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An increase in longitudinal compression force with braiding angle can be clearly observed for most of the stents. This was attributed to the better cover of the stent

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structures at higher braiding angles, resulting in better structural integrity and higher compressive load bearing capability. However, at higher braiding angles, yarns became more oriented towards the transverse axis, thereby reducing their resistance to axial compression. Due to these two opposing factors, for some of the stents there was almost no change in longitudinal compression force with braiding angle.

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Figure 14- Influence of braiding angle on longitudinal compressive force of PA stents

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(a), PES stents (b) and PP stents (c) produced using mandrel diameter of 3.2mm.

4.4. Bending properties

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Bending properties of developed stents were significantly influenced by all studied parameters, namely type of fiber, braiding angle, monofilaments diameter and diameter of mandrel and therefore, it was too difficult to find independent influence of each of

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these parameters on bending properties. However, for samples of PES and PA with lower monofilament diameters (0.27mm. 0.35mm and 0.30mm. 0.40mm) it was observed that an increase in the braiding angle resulted in better bending resistance and higher unchanged diameter. The comparison of bending behaviour of various stents is presented in Figure 15. It is clear that most of the samples retained more than 75% of their original diameter during the bending test (indicated by the straight line).

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Figure 15-Comparison of unchanged diameter of PA stents (a), PES stents (b) and PP

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stents (c).

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4.5. Discussion

Different requirements have been reported in the literature for obtaining optimal stent design. Most of these properties are associated with the mechanical and physical

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properties of the stent’s materials and structures [17]. An ideal stent should have biocompatibility, X-ray and MRI visibility, radial strength, acute and chronic recoil,

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axial and radial flexibility, deliverability, profile, and long term integrity [17]. Most of the above properties are dependent on each other and their careful control is required to obtain an optimal design. For the commercial self-expandable Nitinol stents, the reported important physical and mechanical properties include radial compression, longitudinal compression, bending properties and porosity [15]. Therefore, in order to design a fibrous stent that can replace commercial Nitinol stents, the present research attempted to investigate important material and structural parameters of braided stents and to select suitable parameters to obtain above properties matching with those of Nitinol stents.

The experimental results presented in the previous sections show that the various properties of braided stents are highly dependent on different material and structural factors. Fibre diameter is an important material property which significantly influenced

ACCEPTED MANUSCRIPT all the investigated properties. Higher fibre diameter led to higher radial and axial compression resistance, however, at the cost of reduced porosity. Similar effect was also observed in case of braiding angle. Higher braiding angle improved the mechanical

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performance, but reduced the porosity. On the contrary, the influence of mandrel diameter (i.e. stent diameter) was exactly opposite, i.e. higher stent diameter improved

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the porosity and reduced the mechanical performance. The influence of fibre type was primarily associated with the fiber’s inherent mechanical properties such as elastic modulus, bending rigidity, etc. Among the studied fibres, polyester fibre possesses better mechanical properties and therefore, improved the mechanical performance of the

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stents. Most of the produced stents could maintain more than 75% of their original diameters during bending test, fulfilling the requirement of commercial Nitinol stents.

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Besides these parameters, fractional properties of fibres also had significant influence. For braided stents, fibres should not have very low frictional co-efficient as it results in relative sliding between the fibres and therefore, results in lower resistance towards

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axial or radial compression.

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The fibres used in the present study, i.e. polyester, polyamide and polypropylene have been widely used in medical applications like sutures [8]. These fibres are suitable for

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these applications due to their good mechanical performance. Biocompatibility of implantable medical devices is also an important property and can be strongly influenced by the type of fibres. Overall, the fibres studied in this research have good

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biocompatibility as they lead to low tissue reactions and polypropylene is the best fibre in this aspect due to its highly inert nature.

4.5.1 Selection of fibre type and design parameters The main objective of the present study was to develop 100% fibrous stents possessing similar physical and mechanical properties as commercial stents. Therefore, suitable design parameters have been selected in order to develop stents with properties matching with the commercial Nitinol stents. However, according to the requirements of specific applications, completely different set of parameters can also be selected. In the present case, samples with higher mandrel diameters (5mm and 6mm) and braiding angle (35°) can be selected to produce the best design. Although the increase in mandrel diameter reduced the mechanical performance, it increased the porosity and therefore, higher mandrel diameter was found favorable to obtain a good balance between

ACCEPTED MANUSCRIPT mechanical performance and porosity. Similarly, polyester stents were found more advantageous due to higher mechanical performance as well as good porosity. Moreover, stents produced with lower monofilament diameters can be considered better

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due to higher porosity with sufficient mechanical performance. Considering all these factors, the best specimen produced in this research was PES270A35D6, as the

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properties offered by these stents closely matched with the Nitinol stents. Similar approach of selecting design parameters was also followed by Kim et al. in case of shape memory polyurethane braided stents and in that case, fibre diameter and braiding angle were selected as the design parameters to obtain properties similar to Nitinol

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stents [16].

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4.5.2 Comparison with commercial stents

Commercial Nitinol stents were tested elsewhere [15] and their mechanical properties have been compared with the selected fibrous stent produced within this study

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(PES270A35D6). From Table 3, it is possible to conclude that the developed fibrous stents fulfill the requirements of commercial stents, as far as the physical and

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practical application.

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mechanical properties are concerned and therefore, they can be good candidates for

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Table 3- Comparison of properties between produced fibrous stents and commercial

Properties

Force for Longitudinal compression Force for Radial compression Porosity Unchanged bending diameter

Nitinol stent. Nitinol Stent

PES270A35D6

[15] [16]

Stent

[0.16-5.28]N

0.23N

[1.13-2.9]N

1.29N

>70%

80%

>75%

85.5%

There are some important aspects related to the produced braided stents that should be considered for future research and developments. In the present study, the influence of

ACCEPTED MANUSCRIPT various parameters have been studied experimentally and based on this, best design parameters have been selected to mimic the properties of metallic stents. However, numerical modeling techniques such as finite element method can be used further to

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better study the influence of these parameters and to find the optimum solutions, as performed in some recent studies with Nitinol and polymeric stents [11, 18, 19].

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Deployment process of the developed stents inside blood vessels after their implantation should be further studied, as this aspect is very important regarding the practical application of the produced stents. Fatigue performance of the developed stents should also be evaluated further, as the stents need to withstand continuous pulsatile blood

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flow. Moreover, biocompatibility issues of the developed fibrous stents should be characterized. Final optimization of design parameters will also strongly depend on

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these properties as well as on the type of application. This research study, however, will help to understand the influence of important design factors related to material, structure and production process of braided stents, and can be utilized in future research and

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developments on optimum stent designing based on fibrous materials.

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5. Conclusions

Within this work, 100% fibrous stents were developed using braiding technology and

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the influence of different material, process and structural parameters such as fibre type, braiding angle and mandrel diameter on their performance was thoroughly investigated. According to the experimental results, the best performance (i.e properties closely

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matched with those of commercial metallic stents) was achieved for polyester stents with 0.27 mm yarn diameter, braiding angle of 35⁰ and mandrel diameter of 6mm, providing the values of radial compression, longitudinal compression, porosity and unchanged bending diameter of 1.29 N, 0.23N, 80% and 85.5%, respectively. As the properties of these 100% fibrous stents are in the range offered by commercial metallic stents, they have good potential for commercial use. The results obtained in the present study can be used to understand different parameters that influence the properties of braided stents and will help to design a 100% fibrous stent with targeted set of properties. The use of numerical modeling technique such finite element method will help further to select optimum design parameters, and important issues related to application of produced stents such as deployment process, biocompatibility, etc. should also be studied in future.

ACCEPTED MANUSCRIPT Acknowledgments This project is supported by Portuguese Foundation for Science and Technology (FCT)

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through the PhD grant SFRH / BD / 90321 / 2012.

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References

[1] M. Nichols, N. Townsend, R. Luengo-Fernandez, L. Gray, P. Scarborough e M. Rayner, “European Cardiovascular Disease Statistics 2012. European Heart Network.,” European Society of Cardiology. , 2012.

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[2] A. Freitas, “ Desenvolvimento de stents em fibra de polipropileno para implantação arterial.,” Master Thesis, University of Minho., 2007.

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[3] N. Vila, “Braided Hybrid Stents Design,” em Master Thesis on Design and Marketing, University of Minho, 2009. [4] L. França e A. Pereira, “Atualização sobre endopróteses vasculares (stents): dos estudos experimentais à prática clínica,” Jornal Vascular Brasileiro, 7, 2008.

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[5] S. Irsale e S. Adanur, “ Design and Characterization of Polymeric Stents.,” Journal of Industrial Textiles, 35., pp. 189-200, 2006.

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[6] A. Freitas, M. Araújo, R. Fangueiro e W. Zu, “Freitas, A., Araujo, M., FDevelopment of weft-knitted and braided polypropylene stents for arterial implant.,” Journal of Textile Institute, 101., p. 1027–1034, 2010.

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[7] J. Cruz, S. Rana, R. Fangueiro e R. Guedes, “Designing artificial anterior cruciate ligaments based on novel fibrous structures,” Fibers and Polymers, vol. 15, nº 1, pp. 181186, 2014.

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[8] W. Zhong, “Applications of Braided Structures in Medical Fields,” em Braided Structures and Composites: Production, Properties, Mechanics, and Technical Applications, CRC Press, 2015. [9] S. Zhao, X. Liu e L. Gu, “The Impact of Wire Stent Fabrication Technique on the Performance of Stent Placement,” Journal of Medical Devices, vol. 6, pp. 011007-1 011007-4, 2012. [10] Y. Hoi, C. Ionitaa, R. Tranquebara, K. Hoffmanna, S. Woodwarda, D. Taulbeed, H. Menga e S. Rudina, “MengFlow modification in canine intracranial aneurysm model by an asymmetric stent: studies using DSA and image-based computational fluid dynamics.,” em Proceedings of the Society of Photo-Optical Instrumentation Engineers, 13., 2006. [11] J. H. Kim, T. J. Kang e W.-R. Yu, “Simulation ofmechanical behavior of temperatureresponsive braided stents made o fshape memory polyurethanes,” Journal ofBiomechanics, vol. 43, p. 632–643, 2010. [12] M. Araújo, R. Fangueiro e H. Hong, Têxteis Técnicos- Materiais do novo milénio, Williams, Ltd., 2000.

ACCEPTED MANUSCRIPT [13] P. Potluri, A. Rawal, M. Rivaldi e I. Porat, “Potluri, P., Rawal, AGeometrical modelling and control of a triaxial braiding machine for producing 3D preforms.,” Composites, Part A: applied science and manufacturing, 34. , pp. 481-492, 2003.

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[14] M. Yuksekkaya, “Analysis Of Elstic Deformation Of Braided Tubular Structures for Medical Applications,” Journal of Engineering Sciences, 7, pp. 277-285, 2001.

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[15] Goremedical;, Mechanical Properties of Nitinol Stents and Stent-Grafts: Comparison of 6mm Diameter Devices., L. Gore & Associates, Inc., 2011. [16] J. Kim, T. Kang e W. Yu, “Mechanical modeling of self-expandable stent fabricated using braiding technology.,” Journal of Biomechanics, 41., pp. 3202-3212, 2008.

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[17] P. Poncin e J. Proft, “Stent Tubing: Understanding the Desired Attributes,” em Materials & Processes for Medical Devices Conference, 2003. [18] M. Azaouzi, A. Makradi, e S. Belouettar, “Deployment of a self-expanding stent inside an artery: A finite element analysis,” Materials and Design, vol. 41, p. 410–420, 2012.

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[19] M. Azaouzia, N. Lebaalb, A. Makradia e S. Belouettara, “Optimization based simulation of self-expanding Nitinol stent,” Materials & Design, vol. 50, pp. 917-928, 2013.

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Graphical Abstract

ACCEPTED MANUSCRIPT Highlights 

100% fibrous stents were produced from PES, PA and PP monofilaments using braiding



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technology. Influence of various parameters was studied and the stents with optimum design were



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developed.

Developed stents presented good mechanical properties, porosity and dimensional stability.

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Developed 100% fibrous stents presented similar characteristics to commercial metallic stents.