Manufacturing of stents: Optimize the stent

Transcription

Manufacturing of stents: Optimize the stent with new manufacturing technologies
12
MANUFACTURING OF STENTS:
OPTIMIZE THE STENT WITH NEW
MANUFACTURING TECHNOLOGIES
Andreas SCHUESSLER, Ullrich BAYER, Gerd SIEKMEYER,
Rainer STEEGMUELLER, Martin STROBEL, Admedes SCHUESSLER
We proposed a review of the different driving forces for further development of stent
manufacturing technologies. Influences from the market, the clinics and the customers lead to
manifold research and engineering activities.
We focused on developments in the area of materials development, including biodegradable
stent materials, and on developments and optimization of manufacturing and surface
modification technologies.
1
INTRODUCTION
Since the invention of the balloon expandable stent by Palmaz in 1988, and FDA
approval for peripheral vascular in 1990 and
for coronary vascular procedures in 1994,
more than 10 million people in the US have
undergone a coronary or peripheral stenting
procedure. Such a unique and rapid development in this field of medicine in only 12
years was based on the significantly better
results obtained for stenting compared to
standard balloon angioplasty. Angioplasty
was successful in reopening clogged arteries, however in many cases, immediate reclosure or arterial recoil caused failure of
the angioplasty procedure. Stents provide a
mechanical means of resisting recoil of the
artery after balloon angioplasty, but do not
prevent (and may to a certain extent even
cause) another failure mode known as restenosis. Restenosis is a reclosure of the
treatment site due to cell proliferation caused by local injury and a subsequent tissue
response. With increasing clinical application of and experiences in stenting and
increasing production quantities, requirements for stents and their manufacturing technologies have changed, too.
FACTORS IN OPTIMIZING
STENT TECHNOLOGIES
We divided the driving forces for technology development in the area of stents into
three basic groups: design requirements,
material related properties, and market
aspects as demonstrated in Figure 1.
These aspects are affecting and will continue to affect development in manufacturing technologies in different ways and at
different levels. The details will be discussed in the following chapters
Market
Design
Price
Miniaturization
Clinical
Requirements
Visualisation and
diagnostic
techniques
Customized
surfaces
Radiopacity
Material
Manufacturing
Technologies
Material
combinations
Fatigue
resistance
Tolerances
Drug Adhesion
Biodegradability
MRI visibility
Figure 1: Influences on manufacturing technologies
for stents.
MARKET INFLUENCES CELL
CULTURE TRIALS
The primary market driver for stents has
been the relative ease that most vessels in
the body accept a foreign body such as a
stent implanted on a permanent basis. Ongoing optimizations of the procedure
continue however:
a) prevention of acute and/or late restenosis within the stent via drug
coatings on the stent surface;
b) reduction of the profile of the device
to the smallest possible size in order to
minimize the invasiveness of the
procedure and allow access to even
the most distal locations;
c) increase in stent durability in harsh
environmental anatomic locations
such as the superficial femoral artery
(SFA) where torsion, radial and
longitudinal tension and compression
combine to create material fatigue
challenges;
d) creation of a device that can be seen
easily under x-ray visualization techniques (also known as fluoroscopy);
2
e) creation of a device that can be visualized via non-invasive techniques such
as magnetic resonance imaging (MRI).
Regulatory bodies continue to try to en-sure
safer and more efficacious devices by requesting larger clinical studies with longer
term follow-up. Due to the large amount of
devices recalled either voluntarily or by regulatory injunction from the market in the last
five years1. The Federal Drug Administration
(FDA) is announcing fundamental changes in
the way the agency will track medical devices. Recently, adverse effects (late thrombosis with deadly outcomes) were reported for
long-term clinical studies for drug eluting
stents (CYPHER stent of Cordis and TAXUS
stent of Boston Scientific). Also, broken
struts and broken stents in peripheral arteries,
especially in the SFA are a concern of the
FDA. The FDA has undertaken different
means to tighten the regulatory requirements
related to stents. Among those are the decrease of off-label uses and potentially the
removal of 510 k approvals (a notification of
intent to market) in favour of a PMA process
(a request to market) and the application of
21 CFR part 820 QSR for contract manufacturers.
Finally, cost containment has been an ongoing struggle in health care for the last
decade and with the aging population, this
will continue to be a major economic drain
on society.
MATERIAL CHOICE AND
FUNCTION
Conventional stent materials
Stainless Steel, Tantalum, Niobium, CoCr
alloys and Nitinol have been introduced as
materials for stents in the past 15 years.
However, unmet medical needs and general
technological advancements have led to additional efforts to improve materials properties
and to develop new materials for vascular
stents, especially biodegradable materials
such as magnesium and degradable polymers.
Biodegradable materials
Polymer materials like poly (l-Lactid acid)
PLLA seem to be generally applicable for
drug eluting stents and if they finally make
their way into a commercial application may
offer cost advantages in stent manufacturing
compared to the conventional manufacturing
processes of today’s stents. As an alternative
to polymers, the use of degradable metallic
materials (Figure 2), such as magnesium
alloys and iron, have been investigated during the past 5 years. In general, degradable
materials can be used for indications which
do not require a permanent mechanical enlargement of the artery. However, the main
advantage of biodegradable materials is
assumed to be that any long term risk associated with a foreign object implanted in the
body is obviated and no long term mechanical adverse effects of conventional stents,
such as stent “crushes” or strut fractures2 – to
what extent the degradation process and
degradation products might effect long term
patency is unclear and subject to investigations As such, results on reduction of restenosis rate of biodegradable stents
(magnesium) are inconsistent, ranging from
a significant reduction3 to no difference
compared with balloon angioplasty4.
However, one of the main challenges in
the development of Mg stents is the control
of the degradation mechanism and the lack
of x-ray visibility5,6 (Figure 3) and is therefore subject of intensive investigation7,8,9,10.
3
J&J
Palmaz: SS
Wiktor:
Tantalum
Medtronic
Inflow Dynamics Driver: MP35N
Starflex: Niobium (Cobalt Alloy)
1990´s
Schneider
Wallstent:
Elgiloy*
Biotronik
AMS: Mg
2005
Angiomed
Memotherm:
Nitinol
Angiodynamics
AngioStent: Pt Ir
Guidant
Vision: L605*
Abbott
TriMaxx:
SS/Ta
Composite
Figure 2: Stent Material Introduction
Biopolymers
Figure 3: Magnesium stent (top) versus stainless
steel stent (bottom) approximately 30 days after
implantation in porcine coronaries5.
Biodegradable polymers possess a variety
of challenges, such as radial force, recoil,
large profile, degradation rate and products,
biocompatibility and radiopacity. However,
there are an increasing number of development programs with polymer stents11:
a. The Igaki-Tamai stent is made from
PLLA and features a zig-zag design
which is deployed using a heated
balloon
b. The Reva Medical stent uses tyrosinederived poly (DTE carbonate). This
stent uses ratchet links in order to
achieve a higher radial force and is
loaded with iodine for inherent radiopacity
c. The BVS stent (Bioabsorbable Vascular Solutions, Guidant) is made from
PLLA loaded with the drug everolimus and features both a self expandable and a balloon expandable design
Table I shows a comparison of different
stent material groups with regard to technical and market aspects.
4
Table I. Properties and trends of different stent materials
Remarks
Polymer stents
Metallic stents
Degradable
Degradable
Non-degradable
(non degradable not
considered herein)
(Mg-, Fe-alloys)
(Nitinol, SS, CoCr, Ta)
n/a
n/a
++++
++
++
+
+
+
+++
↑
↑
→
Manufacturing by
Casting, injection
moulding
Laser cutting
Laser cutting, Wire
technologies
Price trends
None marketed yet
Mg↑
SS↓, CoCr→, Nitinol →
Degradability
Drug release
Degradability
No inflammatory
effects compared to
polymers
Proven biocompatibility
Radiopacity
High radial force (SS, CoCr)
Superelasticity and excellent
wall apposition (Nitinol)
Low radiopacity
Degradation
uncontrolled
Fatigue issues
No potential as drug carrier
--
-
++
↑↑
↑↑
→
Current market
value
Expected growth
rate
Clinical experience
Trends in clinical
acceptance
Main advantages
Recoil
Local inflammation
Main disadvantages
Poor vizualisation
Poor mechanical properties
Mechanical
properties
Need for further
research
Radiopacity and MRI visibility
For exact positioning and release at the
treatment site, good visibility of the stent
under fluoroscopy is required. Due to materials of high atomic number conventional
metal stents have significantly better fluoroscopic visibility compared to magnesium or
biodegradable polymer stents (Figure 4).
Nevertheless, some indications, especially
for peripheral stents require an improvement
of the radiopacity by the addition of markers
from tantalum or gold at the end of stents by
riveting or welding12. So far, new concepts
to improve the radiopacity of the less visible
degradable materials are lacking.
The radiopacity can also be improved by
direct alloying elements or ion implantation
with a higher atomic number into the material. For Nitinol the addition of ternary elements like platinum or palladium which still
retains the mechanical properties of the base
material are described in literature13,14.
Figure 5 shows another new technological
solution to enhance radiopacity.
5
Figure 4: Similar stent designs manufactured from
different materials.
By welding or laminating a sandwich
material (e.g. Trimaxx manufactured by
Abott Vascular, stainless steel / tantalum /
stainless steel) radiopacity of a small stent
can be improved appreciably.
Figure 5 : Trimaxx Tube concept of Abott Vascular
Perclose, Inc.
Material interaction with diagnostic clinical visualization tools is important during
MRI. The magnetic field might cause heat
and forces on the stent component due to its
ferromagnetic material behaviour. As the
majority of today’s stents are made out of
nonmagnetic materials, most stents are MRI
safe. However, a stent acts as a Faraday
cage (it shields radio-frequency signals) and
creates a large image artefact. (Figure 6)
shows a standard vascular stent imaged
under MRI.
Figure 6 : Stent images during regular MRI investigations (Source: ©Biophan Technologies Inc.).
In contrast, new surface coatings implementing resonator technology, which uses
tuned circuits to increase the radio frequency (RF) signal, make it possible to image
within and around a stent (Figure 7).
Figure 7 : Improved stent MRI visibility by coating
the metallic stent component (Source: © Biophan
Technologies Inc.).
Customized surfaces
Surface micro patterning
Standard drug eluting stent (DES) technology places active agents which are antagonists of restenosis into a polymer matrix.
Drugs are released into the surrounding
tissue by a controlled degradation process
(Figure 8).
6
Figure 10 : (a) Stent structure with micro structure,
(b) porous surface from oxides, (c) porous surface
from metal
Figure 8 : Drug release through polymer drug pads
which are embedded e.g. by laser cutting cavities
(Source: Conor Medsystems Inc.).
Today, most DES are directly drug coated,
however, recent clinical studies have shown
an increased risk of late thrombosis which is
attributed to this solution. Newer technology
routes use small cavities as reservoirs within
stent struts (Figure 8) or in cavities along
the strut as shown in Figure 9 below.
Figure 9: Cavities to hold drugs by surface structuring (Source: Sorin Biomedical)
In contrast to the above discussed routes
new technology to manufacture porous surfaces (as displayed in Figure 10) are possible, too. Porous surfaces can be built up
from oxides or metallic materials.
Another way to get functionalized surfaces
with a higher medical performance is described in the US Patent 6,190,404. It refers to a
special surface topography of an intravascular stent. The morphology is shown in
(Figure 11).
The grooves are suggested to promote the
migration of endothelial cells onto the stent
surface. Different methods to produce surface morphologies for improved endothelialization such as chemical etching or etching
by laser treatment have been described (US
Patent 2002/0017503).
Figure 11 : Cross section of a stent with a grooved
inner surface (US Patent 6,190,404).
Fracture and fatigue resistance
Fracture and fatigue of vascular stents,
such as coronary stents, AAA stent grafts
and lately SFA stents has been a concern for
a while15-19. While failures in coronary
stents and AAA stent grafts seem to be
highly attributed to high cycle fatigue, most
7
showing that different tube material (different ingot suppliers and tube manufacturers)
result in different lifetime of test samples
(Figure 12).
Strain vs. Number of Cycles Diagram
6
5
Strain [%]
stent fractures in superficial femoral arteries
occur in relatively short time periods and are
due to the high strains/stresses applied to the
implants due to the pronounced mechanical
loading (extension, compression, torsion) in
the limb.
The fracture and fatigue behaviour has
been mainly addressed by design improvements. However, it is well known and
accepted that stent manufacturing processes
do also affect the fatigue behaviour. Excessive heat input during laser cutting can
cause micro cracking and needs to be avoided. Consequently, working with small laser
spot sizes and short pulse duration which
lead to smaller heat affected zones is a
preferred method. It is also state-of-the-art
that the heat affected zone should be entirely
removed during the subsequent processing
steps. It has also been recognized during the
past years that electro polishing does not
just produce corrosion resistant biocompatible surfaces, moreover edge rounding and
the removal of micro cracks and surface
imperfections improves the fracture and
fatigue behaviour even further. With the use
of appropriate laser cutting and electro
polishing techniques and the careful adjustment of all process steps the inducement of
flaws leading to early fracture and fatigue
can be avoided.
Stent material and raw material processing
is another parameter affecting the fatigue
behaviour and is unfortunately not very well
understood. This is especially true for Nitinol due to its complex microstructure and
material processing cycle. However, there is
a number of results suggesting that at least
at high alternating strains (low cycle fatigue)
- which is the predominant load for self
expanding stents in the SFA - is associated
with TiC and/or TiNiO inclusions in the raw
material which act as a crack initiator20-22.
This is supported by our own investigations
A
4
B
3
C
D
2
E
1
0
1000
10000
100000
1000000
Cycles to Failure [10E3]
Figure 12 : Comparison of the fatigue behaviour of
different Nitinol tubing materials determined on
MicroBone (Tensile Test) samples from tube OD:
2.3mm, wall:0.24 tubing at 20°C and a cycling
frequency of 12Hz
MANUFACTURING
TECHNOLOGIES
Laser machining
Solid state lasers (Nd:YAG) using flash
lamps is the standard beam source for stent
cutting. These lasers have been optimised
throughout the past 15 years for precision
cutting and for approximately the past 7
years even special stent cutting laser machines have been made commercially available. For stent cutting of SS, CoCr or Nitinol
stents aspect ratios (Cutting curve width /
wall thickness) of about 12:1 can be achieved. The major disadvantage of those lasers
is the inherent formation of a temperature
gradient within the laser rod during operation. This causes lensing effects and mechanical stresses in the rod leading to a disturbed beam shape with a lower beam quality.
The trend towards device miniaturization
and tighter tolerances for stents requires
smaller cuts and less heat input and con-
8
sequently improved beam quality and pulse
durations. In recent years, new promising
laser resonator architectures have been
developed. In fact, diode-pumped solid-state
lasers (DPSSL) provide a sharper power
distribution i.e. an improved beam profile
quality. Nevertheless, mechanical resonator
instability effects still can affect the cutting
process. Fiber lasers appear to present the
solution for this issue. Housed within a glass
fiber OD <0.5mm, no misalignment can
occur; an optimized beam profile is ensured.
Meanwhile, serviceable laser diodes ready
for direct fiber pumping with short powerup rise times provide high peak powers for
good cutting quality. However, the industriial capability and reliability needs to be
demonstrated.
Although optimized melt cutting lasers are
available, the achievable cutting width for
the final product is still limited downwards.
Since laser melt cutting is a thermal process,
the formation of a heat affected zone (HAZ)
is not avoidable. The enlargement of the
cutting width related to the removal of the
HAZ leads for a given wall thickness
~300µm to a reduced achievable aspect ratio
for the final product of 12:1. In the case of
inert gases used for the process (which is
preferable for fine cutting of Nitinol), burrs
and residual parts remain within the tubing
and the cut structure and need to be removed
semi-automatically. As a result of mechanical impact, a deformation of the part can
occur. Also, melt removal by a gas stream
momentum is strongly restricted to small
cutting widths since an appropriate steam
channel is not established below ~8µm. The
fineness of possible structures is therefore
limited (Figure 13).
Figure 13 : SEM image of laser cut Admedes logo in
Nitinol tubing (top) and the 10µm cutting width of
the “H” letter (bottom).
Ultrashort-pulse lasers offer another way
of cutting materials. No melt is produced as
the laser transfers the material directly from
the solid to the vapour phase. Therefore, no
burrs or slag are produced leading to a clean
surface immediately. Also, due to the very
short energy impact, the whole irradiated
material is removed by vaporisation and
HAZ is not formed. The lateral resolution is
therefore diffraction-limited leading to the
ultimate miniaturization. However, beam
cauterization reduces the possible miniaturization, as the depth of focus leads to an
enlargement of the cutting width at the
upper and lower slot end (Figure 14).
9
Figure 15: Spire stent of Vascular Architects. The
frame is manufactured by photo chemical etching of
Nitinol sheet (left) covered with ePTFE material
(right).
Thin film technology
Figure 14 : SEM images of a resorbable polymer
stent machined with a femto laser system (top) and
magnification of the cutting edge (bottom).
Photo-chemical etching
Photo-chemical etching or photo-chemical
machining has been used to manufacture
stainless steel stents in combination with
forming and welding (NIR stent of Medinol)
or Nitinol stents in combination with shape
setting (aSpire stent of Vascular Architects).
Although photo-chemical etching possesses
a number of advantages, such as stress- and
burr-free machining, suitability for all conventional stent materials and broad use of
the process has not seen (Figure 15). One of
the main disadvantages is the low aspect
ratio depth/width of about 1:1 witch prevents to apply the technique to stent designs
of standard coronary and peripheral stents
(aspect ratio ~ 5:1 to 12:1).
Manufacturing of three dimensional structures of Nitinol for the use as stents, stent
grafts and related devices by thin film (sputter) technology followed by photoetching
have been described for example by Gupka
et al.23 and Banas et al.24. The proposed manufacturing techniques include planar sputter deposition methods as well as sputter deposition on a rotating 3D substrate. Manufacturing of NiTi thin films, first investigated by the Japanese in the late 80s (JP
03162540A) has to be done in an ultra high
vacuum (UHV) atmosphere and is not a
simple technique. Consequently, there have
been only a few research groups worldwide
which were able to achieve promising
results on the functional properties of thin
film nitinol structures. The sputter deposition method for producing medical devices
such as stents has the advantage of making
very thin devices (in the range of a few micrometers) with small features and tight tolerances (Figure 16). Sputter deposited NiTi
films are extremely pure containing no contaminants and are claimed to be less throm-
10
bogenic compared to conventionally manufactured material25. Although sputter deposition seems to have distinct advantages for
niche applications (e.g. devices for very
small vessels), high manufacturing costs and
an unfavourable patent situation seem to
hamper a wide penetration of this technology (Figure 17).
Figure 16: Thin film cylinder manufactured by
sputter deposition15
Figure 17 : Stent Graft consisting out of a fine NiTi
mesh produced by sputter deposition attached to a
NiTi stent conventionally manufactured16
Electroforming
Manufacturing of high aspect ratio micro
metal parts using electroforming involves
photolithography to form the structural information (image) in a photoresist and electroplating of the metal at the openings in the
resist. This process, for example in the form
of the LIGA-technique, is a standard process
in MEMS technology. The company ESI Electroformed Stents Inc., has developed an
electroforming process to produce stents
from Gold, Silver or Nickel. The ESI process coats a conductive mandrel with a photoresist. The resist is exposed to a UV light
source to form the stent image on the cylinder. Developing and stripping the resist forms
the openings in which the metal is deposited. The photoresist and mandrel are dissolved, leaving the electroformed stent. (Source:
www.estent.com) Electroformed stents can
be manufactured very precisely and, if done
in a batch process, very cost effectively. The
disadvantage, however is that it is has not
been achieved applicability of components
out of stainless steel, cobalt-chromium alloys or nitinol. A commercial use of electroformed stents is not known.
CONCLUSIONS
At a first glance, requirements from the
market to improve treatment options with
stents will be more likely achieved by new
and improved stent materials rather then by
the optimisation and development of new
manufacturing processes. In case degradable
materials will start to penetrate the market
new manufacturing technologies will be
required and developed.
Concurrently, the potential of the existing
materials, manufacturing processes and surface modification technologies is not yet
fully optimized. In particular advanced laser
technology and surface patterning and modification technologies (nano technologies)
are still in development.
We expect an incremental improvement of
stents by manufacturing technologies rather
than drastic changes in the next few years.
To achieve this, improved interaction between process engineers, materials and surface scientist and clinical experts is required.
11
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Manufacturing of stents: Optimize the stent

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