U.S. patent application number 15/569670 was filed with the patent office on 2018-05-03 for 3d printing of biomedical implants.
The applicant listed for this patent is Northwestern University. Invention is credited to Guillermo A. Ameer, Evan C. Baker, Cheng Sun, Robert Van Lith, Henry O. T. Ware, Jian Yang, Fan Zhou.
Application Number | 20180117219 15/569670 |
Document ID | / |
Family ID | 57199627 |
Filed Date | 2018-05-03 |
United States Patent
Application |
20180117219 |
Kind Code |
A1 |
Yang; Jian ; et al. |
May 3, 2018 |
3D PRINTING OF BIOMEDICAL IMPLANTS
Abstract
Provided herein are methods, compositions, devices, and systems
for the 3D printing of biomedical implants. In particular, methods
and systems are provided for 3D printing of biomedical devices
(e.g., endovascular stents) using photo-curable biomaterial inks
(e.g., or methacrylated poly(diol citrate)).
Inventors: |
Yang; Jian; (Evanston,
IL) ; Baker; Evan C.; (Chicago, IL) ; Ware;
Henry O. T.; (Shawnee, OK) ; Zhou; Fan;
(Torrance, CA) ; Sun; Cheng; (Willmette, IL)
; Ameer; Guillermo A.; (Chicago, IL) ; Van Lith;
Robert; (Evanston, IL) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Northwestern University |
Evanston |
IL |
US |
|
|
Family ID: |
57199627 |
Appl. No.: |
15/569670 |
Filed: |
April 28, 2016 |
PCT Filed: |
April 28, 2016 |
PCT NO: |
PCT/US16/29774 |
371 Date: |
October 26, 2017 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
62154499 |
Apr 29, 2015 |
|
|
|
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61F 2/82 20130101; B29C
41/00 20130101; B29K 2995/006 20130101; B33Y 70/00 20141201; B29C
64/20 20170801; C09D 167/06 20130101; B33Y 10/00 20141201; C08J
2367/06 20130101; B29C 64/135 20170801; A61L 31/16 20130101; A61L
31/18 20130101; B29K 2105/0005 20130101; A61F 2240/001 20130101;
B33Y 80/00 20141201; C08J 3/24 20130101; B29C 41/22 20130101; B29L
2031/7534 20130101; B29K 2995/0056 20130101; B33Y 30/00 20141201;
A61L 31/06 20130101; B29K 2033/00 20130101 |
International
Class: |
A61L 31/06 20060101
A61L031/06; B33Y 30/00 20060101 B33Y030/00; B29C 64/135 20060101
B29C064/135; B29C 64/20 20060101 B29C064/20; B33Y 80/00 20060101
B33Y080/00; B33Y 70/00 20060101 B33Y070/00; A61L 31/18 20060101
A61L031/18; A61L 31/16 20060101 A61L031/16; C08J 3/24 20060101
C08J003/24; C09D 167/06 20060101 C09D167/06 |
Claims
1. A system comprising: (a) a photo-curable biomaterial ink; and
(b) a 3D printing device for: (i) dispensing a layer of the
photo-curable biomaterial ink in a pattern according to encoded
instructions, (ii) exposing the layer of the photo-curable
biomaterial ink to light to cure the biomaterial ink and produce a
solidified biomaterial layer, and (iii) repeating steps (i) and
(ii), with each successive layer built upon the previous layer to
produce a 3D structure of the solidified biomaterial.
2. The system of claim 1, wherein the photo-curable biomaterial ink
comprises methacrylated poly(diol citrate).
3. The system of claim 1, wherein the poly(diol citrate) comprises
a polymer of citric acid and HO--(CH.sub.2).sub.n--OH, wherein n is
2-20.
4. The system of claim 1, wherein the photo-curable biomaterial ink
further comprises one or more of: a solvent, a photoinitiator, a
co-initiator, a free-radical quencher, and a UV-absorber.
5. The system of claim 1, wherein the 3d printing device is
configured for laser scanning stereolithography, projection
stereolithography, ink-jet printing, continuous liquid interface
production, or combinations thereof.
6. A biomaterial device produced using a system of one or claims
1-5.
7. A biomaterial ink comprising methacrylated poly (diol citrate),
solvent or dilutant, and a photoinitiator.
8. The biomaterial ink of claim 7, wherein the poly (diol citrate)
is a polymer of citric acid and an aliphatic diol selected from
selected from HO--(CH.sub.2).sub.n--OH, wherein n is 2-20.
9. The biomaterial ink of claim 7, wherein the methacrylated poly
(diol citrate) is present in the biomaterial ink at 50-99 wt %.
10. The biomaterial ink of claim 7, wherein the solvent or dilutant
is present in the biomaterial ink at 1-49.9 wt %.
11. The biomaterial ink of claim 7, wherein the photoinitiator is
present in the biomaterial ink at 0.1-5 wt %.
12. The biomaterial ink of claim 7, further comprising a
co-initiator, a free-radical quencher, and/or a UV-absorber.
13. The biomaterial ink of claim 7, further comprising a
radiopacity agent.
14. The biomaterial ink of claim 13, wherein the radiopacity agent
is selected from the group consisting of: iohexyl, iopromide,
ioversol, ioxaglate and iodixanol
15. The biomaterial ink of claim 7, further comprising a
therapeutic agent.
16. The biomaterial ink of claim 15, wherein the therapeutic agent
is selected from the group consisting of: anticoagulants,
antithrombotic agents, antiplatelet agents, anti-inflammatory
agents, anti-proliferative agents, immunosuppresive agents,
cytostatic drugs, lipid-lowering agents, and antioxidants.
17. A biomaterial device produced by the curing of a biomaterial
ink of one of claims 7-16.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] The present invention claims priority to U.S. Provisional
Patent Application 62/154,373, filed Apr. 29, 2015, which is
incorporated by reference in its entirety.
FIELD
[0002] Provided herein are methods, compositions, devices, and
systems for the 3D printing of biomedical implants. In particular,
methods and systems are provided for 3D printing of biomedical
devices (e.g., endovascular stents) using photo-curable biomaterial
inks (e.g., or methacrylated poly(diol citrate)).
BACKGROUND
[0003] Biodegradable stents (BDSs''), both metallic and polymeric,
offer promising alternatives to conventional bare metal stents
(BMSs) and drug-eluting stents (DESs) in providing temporary drug
release for vessel patency, resisting late stent thrombosis due to
uncovered struts, and potential reduction in the usage of
antiplatelet drugs (refs. 4, 9; incorporated by reference in their
entireties). Moreover, disappearance of BDS over time allows for
eventual recurrence of natural vasomotion.
SUMMARY
[0004] Provided herein are methods, compositions, devices, and
systems for the 3D printing of biomedical implants. In particular,
methods and systems are provided for 3D printing of biomedical
devices (e.g., endovascular stents) using photo-curable biomaterial
inks (e.g., or methacrylated poly(diol citrate)).
[0005] In some embodiments, provided herein are systems comprising:
(a) a photo-curable biomaterial ink; and (b) a 3D printing device
for: (i) dispensing a layer of the photo-curable biomaterial ink in
a pattern according to encoded instructions, (ii) exposing the
layer of the photo-curable biomaterial ink to light to cure the
biomaterial ink and produce a solidified biomaterial layer, and
(iii) repeating steps (i) and (ii), with each successive layer
built upon the previous layer to produce a 3D structure of the
solidified biomaterial. In some embodiments, the photo-curable
biomaterial ink comprises methacrylated poly(diol citrate). In some
embodiments, the poly(diol citrate) comprises a polymer of citric
acid and HO--(CH.sub.2)--OH, wherein n is 2-20. In some
embodiments, the photo-curable biomaterial ink further comprises
one or more of: a solvent, a photoinitiator, a co-initiator, a
free-radical quencher, and a UV-absorber. In some embodiments, the
3D printing device is configured for laser scanning
stereolithography, projection stereolithography, ink-jet printing,
continuous liquid interface production, or combinations
thereof.
[0006] In some embodiments, provided herein is a biomaterial device
produced using a system described herein (e.g., biomaterial ink and
3D printing device).
[0007] In some embodiments, provided herein are biomaterial inks
comprising methacrylated poly (diol citrate), solvent or dilutant,
and a photoinitiator. In some embodiments, the poly (diol citrate)
is a polymer of citric acid and an aliphatic diol selected from
selected from HO--(CH.sub.2).sub.n--OH, wherein n is 2-20. In some
embodiments, the methacrylated poly (diol citrate) is present in
the biomaterial ink at 50-99 wt % (e.g., 50%, 55%, 60%, 65%, 70%,
75%, 80%, 85%, 90%, 95%, 99%, or ranges therebetween). In some
embodiments, the solvent or dilutant is present in the biomaterial
ink at 1-49.9 wt % (e.g., 1%, 2%, 3%, 4%, 5%, 10%, 15%, 20%, 25%,
30%, 35%, 40%, 45%, 49%, 49.9%, or ranges therebetween). In some
embodiments, the photoinitiator is present in the biomaterial ink
at 0.1-5 wt % (e.g., 0.1%, 0.2%, 0.3%, 0.4%, 0.5%, 1%, 2%, 3%, 4%,
5%, or ranges therebetween). In some embodiments, a biomaterial ink
further comprises a co-initiator, a free-radical quencher, and/or a
UV-absorber. In some embodiments, a biomaterial ink further
comprises a radiopacity agent (e.g., iohexyl, iopromide, ioversol,
ioxaglate, iodixanol, etc.). In some embodiments, a biomaterial ink
further comprises a therapeutic agent (e.g., anticoagulant (e.g.,
heparin, Coumadin, protamine, hirudin, etc.), antithrombotic agent
(e.g., clopidogrel, heparin, hirudin, iloprost, etc.), antiplatelet
agent (e.g., aspirin, dipyridamole, etc.), anti-inflammatory agent
(e.g., methylprednisolone, dexamethasone, tranilast, etc.),
anti-proliferative/immunosuppresive agent (e.g., trapidil,
tyrphostin, rapamycin, FK-506, mycophenolic acid), cytostatic drug
(e.g., paclitaxel, rapamycin, rapamycin analogs (e.g., everolimus,
tacrolimus, etc.), etc.), lipid-lowering agent (e.g., statin),
antioxidant (e.g., probucol, vitamin C, retinoids, resveratrol,
etc.)).
[0008] In some embodiments, provided herein is a biomaterial device
produced by the curing of a biomaterial ink described herein.
DEFINITIONS
[0009] Although any methods and materials similar or equivalent to
those described herein can be used in the practice or testing of
embodiments described herein, some preferred methods, compositions,
devices, and materials are described herein. However, before the
present materials and methods are described, it is to be understood
that this invention is not limited to the particular molecules,
compositions, methodologies or protocols herein described, as these
may vary in accordance with routine experimentation and
optimization. It is also to be understood that the terminology used
in the description is for the purpose of describing the particular
versions or embodiments only, and is not intended to limit the
scope of the embodiments described herein.
[0010] Unless otherwise defined, all technical and scientific terms
used herein have the same meaning as commonly understood by one of
ordinary skill in the art to which this invention belongs. However,
in case of conflict, the present specification, including
definitions, will control. Accordingly, in the context of the
embodiments described herein, the following definitions apply.
[0011] As used herein and in the appended claims, the singular
forms "a", "an" and "the" include plural reference unless the
context clearly dictates otherwise. Thus, for example, reference to
"a mPDC polymer" is a reference to one or more mPDC polymers and
equivalents thereof known to those skilled in the art, and so
forth.
[0012] As used herein, the term "comprise" and linguistic
variations thereof denote the presence of recited feature(s),
element(s), method step(s), etc. without the exclusion of the
presence of additional feature(s), element(s), method step(s), etc.
Conversely, the term "consisting of" and linguistic variations
thereof, denotes the presence of recited feature(s), element(s),
method step(s), etc. and excludes any unrecited feature(s),
element(s), method step(s), etc., except for ordinarily-associated
impurities. The phrase "consisting essentially of" denotes the
recited feature(s), element(s), method step(s), etc. and any
additional feature(s), element(s), method step(s), etc. that do not
materially affect the basic nature of the composition, system, or
method. Many embodiments herein are described using open
"comprising" language. Such embodiments encompass multiple closed
"consisting of" and/or "consisting essentially of" embodiments,
which may alternatively be claimed or described using such
language.
[0013] As used herein, the term "polymer" refers to a chain of
repeating structural units or "monomers", typically of large
molecular mass. Examples of polymers include homopolymers (single
type of monomer subunits), copolymers (two types of monomer
subunits), and heteropolymers (e.g., three or more types of monomer
subunits).
[0014] As used herein, the term "linear polymer" refers to a
polymer in which the molecules form long chains without branches or
crosslinked structures.
[0015] As used herein, the term "branched polymer" refers to a
polymer comprising a polymer backbone with one or more additional
monomers, or chains or monomers, extending from polymer backbone.
The degree of interconnectedness of the "branches" is insufficient
to render the polymer insoluble.
[0016] As used herein, the term "pre-polymer" refers to linear or
branched polymers (e.g., not significantly crosslinked) that have
the capacity to be crosslinked under appropriate conditions (e.g.,
to form a thermoset), but have not been subjected to the
appropriate conditions.
[0017] As used herein, the term "crosslinked polymer" refers to a
polymer with a significant degree of interconnectedness between
multiple polymer strands, the result of which is an insoluble
polymer network (e.g., a thermoset). For example, multiple polymer
stands may be crosslinked to each other at points within their
structures, not limited to the ends of the polymer chains. In some
embodiments, two or more different polymers may be crosslinked.
[0018] As used herein, the terms "composite" and "composite
material" refer to materials or compositions generated from the
combination of two or more constituent materials (e.g., compounds,
polymers, etc.). The constituent materials may interact (e.g.,
non-covalently) at the microscopic or molecular level, but
typically do not react chemically (e.g., covalently). At the
macroscopic level, the constituent materials appear homogenous.
[0019] As used herein, the term "biocompatible" refers to
materials, compounds, or compositions means that do not cause or
elicit significant adverse effects when administered to a subject.
Examples of possible adverse effects include, but are not limited
to, excessive inflammation, excessive or adverse immune response,
and toxicity.
[0020] As used herein, the term "biostable" refers to compositions
or materials that do not readily break-down or degrade in a
physiological or similar aqueous environment. Conversely, the term
"biodegradeable" refers herein to compositions or materials that
readily decompose (e.g., depolymerize, hydrolyze, are enzymatically
degraded, disassociate, etc.) in a physiological or other
environment.
[0021] As used herein, the term "subject" broadly refers to any
animal, including but not limited to, human and non-human animals
(e.g., dogs, cats, cows, horses, sheep, poultry, fish, crustaceans,
etc.). As used herein, the term "patient" typically refers to a
subject that is being treated for a disease or condition.
BRIEF DESCRIPTION OF THE DRAWINGS
[0022] FIG. 1. Chemical structure and proton nuclear magnetic
resonance spectrum of methacrylated poly(1,12-dodecanediol citrate)
polymer (left); schematic showing the reaction due to exposure to
UV (right).
[0023] FIG. 2. (a) UV/Vis absorption spectra of Irgacure 819,
Camphorqinone and 2-hydroxy-2-methylpropiophone (Homp) in ethanol;
(b) UV/Vis absorption spectra of Camphorqinone at different
concentrations; (c) Dynamic viscosities of methacrylated
methacrylated poly(1,12-dodecanediol citrate) (mPDC) polymer
solutions with different amount of ethyl acetate; (d) Compression
strength of in-situ mPDC stents of different thicknesses, the stent
is 21.8 mm.times.5.0 mm (length.times.outer diameter).
[0024] FIG. 3. (a) Sketch and gross image of typical repeating
stent element and full 3D CAD Design of the stent; (b & c)
Scanning electron microscopy images of a printed mPDC stent showing
the 20 um layers.
[0025] FIG. 4. (a & c) Low and high magnification of SEM images
of a mesh mPDC-HDDA stent., CAD design shown in the top right
corner; (b & d) low and high magnification of SEM images of
sinusoidal an mPDC-HDDA stent. CAD design shown in the top right
corner.
[0026] FIG. 5. (a) Schematic view of 3-point bending experiment
with a gap of 16 mm and compressive displacement of 3.2 mm for
stent with 5 mm outer diameter; (b) compressive displacement and
resilience of Nitinol (21.8 mm.times.5.0 mm.times.0.2 mm) and mPDC
stent (21.8 mm.times.5.0 mm.times.0.5 mm, (length.times.outer
diameter.times.thickness)); (c) Simulated maximum
force-displacement curves for different thickness stents and
force-thickness curve at onset of kinking in 3-point bending
simulation; (d) Simulated loading and displacement field for 350 um
stent in 3-point bending simulation.
[0027] FIG. 6. (a) Scaled Applied Force vs. Maximum Displacement
profile of stents with 300 um, 350 um, 400 um and 500 um in
thickness. (b) Compressive strength of Nitinol BMS (21.8
mm.times.5.0 mm.times.0.2 mm), HDDA and mPDC printed stents with
9.1 mm.times.5.5 mm in length.times.outer diameter, the thickness
of printed stent is 500 um; (c) Scaled Maximum Usable Applied Force
vs. Stent Thickness curve; (d) Typical Displacement distribution
for Parallel-Plate Compression simulation of stent with 400 um in
thickness.
[0028] FIG. 7. Exemplary process for stent generation by the
methods described herein.
[0029] FIG. 8. Micro-CLIP printing system schematic. UV light is
projected through a thin oxygen permeable membrane. Liquid polymer
material solidifies in the pattern projected and the build platform
raises vertically out of the liquid material bath.
[0030] FIG. 9. Dimensional Differential vs Light Intensity (% of
max). Beside each plot is the corresponding SEM micrograph of the
closest experimentally tested value to dimensional accuracy.
[0031] FIG. 10. Stents Printed with various materials and how that
impacts their aesthetics. Stent #1: 47.79% mPDC, 50% DEF, 0.01%
Sudan I, 2.2% Irg 819. Stent #2: 97.78% HDDA, 0.02% Benzotriazol,
2.2% Irg 819 Stent #3: 97.79% HDDA, 0.01% Sudan I, 2.2% Irg 819.
Stent #4: 80% mPDC, 17% DEF, 2.2% Irg 819, 0.1% Sudan I. Stent #5:
50% DEF, 47.79% mPDC, 0.01% Sudan I, 2.2% Irgacure 819. Stent#6:
50% mPDC, 47.79% DEF, 2.2% Irgacure 819, 0.01% Benzotriazol. Stents
#7-9: 60% mPDC, 35.58% Ethanol, 4.4% Irgacure 651, 0.02%
Benzotriazol.
[0032] FIG. 11. Base Design: (a) CAD Drawing of full length stent.
(b) CAD drawing of unit length of stent. (c) Unit Cell Design. (d)
Scanning electron micrograph of design from CLIP process.
[0033] FIG. 12. Arrowhead Design: (a) CAD of full stent (b) CAD of
unit length of stent (c) Unit cell design (D) scanning electron
micrograph of Arrowhead design after CLIP process.
[0034] FIG. 13. Optimization information: (a) Design variables. (b)
Flexibility test conditions. (c) Contour graph of Flexibility
Metric (FM).
[0035] FIG. 14. Flexibility Optimized Base Design: (a) CAD of full
length stent (b) Unit cell of stent design.
[0036] FIG. 15. Radial compression of 3D-printed stents at
different UV intensities: A) 50% DEF, 47.78% mPDC, 2.2% Irgacure
819 and 0.02% Sudan I, B) 50% DEF, 47.76% mPDC, 2.2% Irgacure 819
and 0.04% Sudan I, C) 50% DEF, 47.72% mPDC, 2.2% Irgacure 819 and
0.08% Sudan I, and D) Radial compressive load at 20% radial
compression for all stents. Black dashed line indicates the target
radial load of a control bare-metal nitinol stent. Stents were
post-cured at 2.times.2.5 minutes.
[0037] FIG. 16. Effect of post-curing time on mechanical strength
of stents. All stents were printed with biomaterial ink of
following composition: 50% DEF, 47.78% mPDC, 2.2% Irgacure 819 and
0.02% Sudan I. UV intensity for printing process was 100%.
[0038] FIG. 17. Relation between stent dimensions and radial
compressive load at 20% radial compression: A) Axial dimension, B)
Laeral dimension and C) Diagonal dimension. All stents were printed
from biomaterial ink of following composition: 50% DEF, 2.2%
Irgacure 819, and mPDC and Sudan I adding up to 47.8% together.
Shaded box indicates the target dimensions, comparable to currently
developed bioresorbable stents. Stents were 3D-printed at various
UV intensities, but all were post-cured at 2.times.2.5 minutes.
[0039] FIG. 18. Mechanical properties of Arrowhead design stents:
A) Dependency on the wall thickness varying between 250-600 um, B)
Dependency on the strut dimensions varying between 150-200 um.
Stents were printed from biomaterial ink of following composition:
50% DEF, 2.2% Irgacure 819, 47.72% mPDC and 0.08% Sudan I. Stents
were postcured at 2.times.2.5 minutes.
[0040] FIG. 19. Mechanical properties of biomaterial ink with added
accelerator compound. Stents were printed from biomaterial ink of
following composition: 50% DEF, 1% Irgacure 819, 47.92% mPDC, 0.08%
SUdan I and 1% Ethyl-4-Dimethylamine Benzoate (EDAB). Stents were
3D-printed at various UV intensities, but were post-cured at
2.times.2.5 minutes.
[0041] FIG. 20(a-e). Temporal series of images showing sheathing
through compression of a 6.5 mm outer diameter stent to 3.1 mm, and
subsequent self-expansion upon sheath retraction of 3D-printed
stent. Full expansion to original diameter reached in approximately
3 minutes.
[0042] FIG. 21. The 3D-printed stents from mPDC are antioxidant,
biocompatible and biodegradable: Left) mPDC scavenges
2,2'-azino-bis(3-ethylbenzothiazoline-6-sulphonic acid) (ABTS) free
radical, Center) Human smooth muscle cells on mPDC show good
spreading and viability. Scale=100 um, and Right) 3D-printed stents
degrade in PBS at 37 C, with approximately 25% degraded after 6
months incubation.
[0043] FIG. 22. Exemplary synthesis of methacrylated poly(diol
citrate).
DETAILED DESCRIPTION
[0044] Provided herein are methods, compositions, devices, and
systems for the 3D printing of biomedical implants. In particular,
methods and systems are provided for 3D printing of biomedical
devices (e.g., endovascular stents) using photo-curable biomaterial
inks (e.g., or methacrylated poly(diol citrate)).
[0045] Among natural and synthetic biodegradable polymers,
chitosan, poly(4-hydroxybutyrate) (PHB),
poly(.epsilon.-caprolactone) (PCL), poly(L-lactide) (PLLA) and
poly(D,L-lactide) (PDLLA) and its copolymers or composites have
been extensively investigated for use in resorbable devices (refs.
10-16; incorporated by reference in their entireties). In
particular, a polylactide stents (e.g., Igaki Tamai or
bioabsorbable vascular stents (BVSs)) have been shown to degrade
into metabolites such as lactic acid, CO.sub.2 and H.sub.2O in two
years and testing indicates they are safe when used in human
coronary arteries (ref. 13; incorporated by reference in its
entirety). Relative to a BMS, the self-expandable PLLA stents
require 8 min for full-expansion in an aqueous environment due to
the viscoelastic behavior of polymer 37.degree. C. (ref. 17;
incorporated by reference in its entirety), which increases the
risk of ischemia and myocardial infarction. Late shrinkage after
degradation also remains a concern. As with metal stents, there are
manufacturing challenges for strut design, processing, and
fabrication.
[0046] Rapid prototyping techniques such as stereolithography,
selective laser sintering, fused deposition modeling and others
have been developed for high precision manufacturing of customized
biomedical devices, greatly expanding in biomedical research and
tissue engineering for a broad range of functional and structural
materials such as hydrogels, polymers and ceramics (refs. 18, 19;
incorporated by reference in their entireties). Continuous tool
path planning strategies have been optimized for open sourced and
commercial fused deposition machines (FDM), making a customized
tracheal stent rapidly and affordably (ref. 20; incorporated by
reference in its entirety). In comparison to selective laser
sintering and solid ground curing, stereolithography offered the
best surface finish in the process of customized tracheobronchial
stents, while selective ground curing had the best repeatability of
length (ref. 21; incorporated by reference in its entirety). Unlike
the above large-size stents, a bioabsorbable drug-coated stent was
manufactured with a 300 um strut diameter using PCL polymer and a
rapid prototyping technique (ref. 22; incorporated by reference in
its entirety). These stents showed to be effective in reducing
neointimal hyperplasia, inflammation and thrombosis formation. A 3D
micro jetting free molding technique has been developed to
fabricate slide or snap fastener biodegradable stents with
polydioxanone (PDO) (ref. 23; incorporated by reference in its
entirety). Different from extrusion techniques in 3D printing,
projection microstereolithography (P.mu.SL) offers a high precision
and high resolution processing method with a digital micromirror
device (DMDTM, Texas Instruments) as a dynamic mask (refs. 24, 25;
incorporated by reference in their entireties)].
[0047] Provided herein are methods for rapid fabrication of
biomedical devices (e.g., implants (e.g., endovascular stents),
etc.) using biomaterial ink and 3D printing or additive
manufacturing processes with micrometer accuracy. Provided herein
are biomaterial inks that are suitable for 3D printing processes,
digital representation of stent design using Computer-aid design
(CAD) modeling, devices (e.g., stents) with optimized mechanical
properties using, for example, numerical simulation, fabrication
processing parameters for device prototype and scalable
manufacturing biomaterial ink that is photopolymerized by
ultraviolet or visible light at various wavelengths, etc. Different
structures (e.g., stent structures), such as sinusoidal formed
wire, helical wrap, and/or laser-fused struts are obtainable and
customizable with patient-specific features in the CAD model and
subsequently fabricated using 3D printing systems with high
fidelity. By optimizing the stent geometry, biomaterial ink
compositions (e.g., polydiolcitrate solution composition),
initiator concentration, and curing conditions, the mechanical
properties of printed devices (e.g., stents) are tailored to
closely match with blood vessel or a bare metal stent. In some
embodiments, kink-resist stents are obtained by incorporating the
stent strut exhibiting near-zero or negative Poisson's ratio. In
some embodiments, the use of biodegradable materials allows for the
encapsulation and slow release of drugs or other agents from the
bulk of the stent rather than a coating that is applied to the
stent struts.
[0048] In some embodiments, using photo-curable polymers, complex
3D microstructures are created. A series of citrate-based polymers
with a wide range of properties such as controllable elasticity,
biodegradability, shape-memory and antioxidant properties have been
developed [26, 27; incorporated by reference in their entireties],
and find use in embodiments herein. After methacrylation with
glycidyl methacrylate, 2-aminoethyl methacrylate, or another
suitbale compound, polymers are printed (e.g., via projection
stereolithography, via Micro-CLIP, etc.) under the appropriate
solvent and additive conditions. Exemplified herein are
compositions and methods to feasibly 3D print complex strut
structures of biodegradable polymers on a micron scale.
[0049] Embodiments herein find use in, for example: endovascular
stents and stent-related implants, 3D printed bio-medical implants
containing patient-specific features, tailoring the mechanical
properties of 3D printed devices through structural and materials
design, related 3D printed products derived from biocompatible
and/or biodegradable biomaterial inks, 3D printed bio-medical
implants for sustained drug release, in vivo sensing platforms,
etc.
[0050] Advantages of some embodiments herein include: the building
materials of the 3D printed stent are precisely tailored to exhibit
a compliant compressive, strength and flexibility with blood vessel
and bare metal stent, the use of biodegradable biomaterial ink
allows for the encapsulation of therapeutic agents, allowing, for
example, the slow release of drugs from the bulk of the device
(e.g., stent) in contrast, to the state-of-the-art coating method
to coat the drug on the surface of stent struts.
[0051] Embodiments herein utilize various 3D printing and/or
additive manufacturing to create biocompatible and biodegradable
devices from compositions comprising biomaterial inks for use, for
example, in various biomedical applications. In some embodiments, a
biomaterial ink comprises a curable (e.g., chemically-curable,
photo-curable, etc.) polymer material. In some embodiments, the
biomaterial ink comprises a polymer component displaying one or
more curable (e.g., chemically-curable, photo-curable, etc.)
substituents; upon exposure of the biomaterial ink to curing
conditions, the biomaterial ink is converted from a pre-polymer
into an insoluble, crosslinked polymeric material.
[0052] In some embodiments, compositions and composites (e.g.,
biomaterial ink and/or solid biomaterials produced therefrom)
described herein comprise a polymeric component. In some
embodiments, a polymeric component comprises a polymer selected
from a polyester, poly(diol citrate) (e.g., poly(butanediol
citrate), poly(hexanediol citrate), poly(octanediol citrate),
poly(decanediol citrate), poly(dodecanediol citrate),
poly(hexadecanediol citrate), etc.), poly(hydroxyvalerate),
poly(lactide-co-glycolide), poly(hydroxybutyrate),
poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,
poly(glycolic acid), poly(glycolide), poly(L-lactic acid),
poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),
poly(caprolactone), poly(trimethylene carbonate), polyester amide,
or co-polymers or composites thereof.
[0053] In some embodiments, a polymeric component comprises a
citric acid-based polymer. In some embodiments, a polymer is the
polyesterification product of one or more acids (e.g., succinic
acid, glutaric acid, adipic acid, pimelic acid, suberic acid,
azelaic acid, sebacic acid, dodecanedioic acid, shorter or longer
linear aliphatic diacids, citric acid, isocitric acid, aconitic
acid, propane-1,2,3-tricarboxylic acid, trimesic acid, itaconic
acid, maleic acid, etc.) and one or more diols or triols (e.g.,
polyethylene glycol, glycerol, linear aliphatic diol (e.g.,
butanediol, hexanediol, octanediol, decanediol, dodecanediol, and
shorter or longer linear aliphatic diols), etc.).
[0054] In some embodiments, a polymer is the polyesterification
product of at least citric acid and one or more linear aliphatic
diols (butanediol, hexanediol, octanediol, decanediol,
dodecanediol, or any linear aliphatic diol from about 2-20 carbons
in length). A polymer may comprise only citric acid and linear
aliphatic diol components or may further comprise additional
monomer components (e.g., sebacic acid, polyethylene glycol,
glycerol, etc.). In some embodiments, a polymer comprises
additional substituents or functional groups appended to the
polymer (e.g., ascorbic acid, glycerol, a NONOate group, etc.).
[0055] In some embodiments, a polymeric component comprises citric
acid as a monomer (e.g., along with a diol monomer). Citric acid is
a reactive tricarboxylic acid that is part of the Krebs cycle and
has been used as a key reactant monomer for the synthesis of
polydiolcitrates with a wide range of properties and uses (Yang,
J., et al., Synthesis and evaluation of poly(diol citrate)
biodegradable elastomers. Biomaterials, 2006. 27(9): p. 1889-98;
U.S. Pat. No. 8,772,437; U.S. Pat. No. 8,758,796; U.S. Pat. No.
8,580,912; U.S. Pat. No. 8,568,765; U.S. Pub. No. 2014/0155516;
U.S. Pub. No. 2014/0135407; herein incorporated by reference in
their entireties). Depending on the diol of choice, materials with
controllable elasticity, biodegradability, and antioxidant
properties can be developed (Serrano et al. Adv Mater, 2011.
23(19): p. 2211-5; Yang J., et al., A thermoresponsive
biodegradable polymer with intrinsic antioxidant properties.
Biomacromolecules, 2014. 15(11):3942-52; U.S. Pub. No.
2014/0037588; herein incorporated by reference in its
entirety).
[0056] In some embodiments, a polymeric component is a poly(diol
citrate), for example, those described in U.S. Pat. No. 8,911,720;
herein incorporated by reference in its entirety. In some
embodiments, derivatives of such poly(diol citrates) are provided.
In some embodiments, a pre-polymer of citric acid and diol is
formed (e.g., by reaction at about 140.degree. C. or other suitable
conditions). In some embodiments, a pre-polymer is reacted with one
or more additional compounds to produce a functionalized (e.g.,
methacrylated) pre-polymer.
[0057] As addressed above, in some embodiments, the curable polymer
component of a biomaterial ink comprises a polymer displaying one
or more curable (e.g., chemically-curable, photo-curable, etc.)
groups. In some embodiments, a curable group is or comprises a
methacrylate or acrylate group. In some embodiments, a curable
group is or comprises N-Vinylpyrrolidone (NVP) or
styrenestyrene.
[0058] In some embodiments, a pre-polymer (e.g., of poly(diol
citrate)) is reacted (e.g., at about 40-100.degree. C.) with a
modifying group to produce a poly(diol citrate) polymer displaying
a curable substituent group (e.g., methacrylate) and/or crosslinked
to form an elastomer displaying the substituent group. In some
embodiments, suitable reactant for modifying the poly(diol citrate)
pre-polymer is glycidyl methacrylate or 2-aminoethyl methacrylate.
In some embodiments, poly(diol citrate) and glycidyl methacrylate
(or 2-aminoethyl methacrylate) are reacted in the presence of
tetrahydrofuran and imidazole. Other substituents (e.g., other than
glycidyl methacrylate) may also be reacted with the poly(diol
citrate) (e.g., alone or with glycidyl methacrylate), and/or other
polymer components may be methacrylated. In some embodiments,
rather than methacrylation, and acrylate group is displayed on the
polymer or pre-polymer to produce a curable polymer for a
biomaterial ink.
[0059] In some embodiments, a citric acid-based, curable polyester
comprises:
##STR00001##
wherein R is selected from H, a poly(diol citrate), and a curable
group (e.g., photo-curable group (e.g., methacrylate group));
wherein R' is selected from H, and a poly(diol citrate); wherein m
is 2 to 20; and wherein at least one R is a curable group (e.g.,
photo-curable group (e.g., methacrylate group)).
[0060] In some embodiments, the citric acid-based polyester
comprises:
##STR00002##
wherein R is selected from H, a poly(diol citrate), and a curable
group (e.g., photo-curable group (e.g., methacrylate group));
wherein R' is selected from H, and a poly(diol citrate); wherein m
is 2 to 20; wherein n is 1 to 1000, and wherein 1-100% (e.g., 1%,
2%, 5%, 10%, 15%, 20%, 30%, 40%, 50%, 60%, 70%, 80%, 90%, 95%, 99%,
100%) of R groups are a curable group (e.g., photo-curable group
(e.g., methacrylate)). In some embodiments, at least one R group
comprises a methacrylate. In some embodiments, the citric
acid-based polyester comprises:
##STR00003##
wherein m is 2 to 20.
[0061] In some embodiments, the citric acid-based polyester
comprises: the citric acid-based polyester comprises:
##STR00004##
wherein m is 2 to 20; and wherein n is 1 to 1000. In some
embodiments, m is 6 to 14.
[0062] In some embodiments, provided herein are methods of
preparing methacrylated poly(diol citrate) comprising: a)
synthesizing a prepolymer of citric acid and an aliphatic diol; and
b) reacting the prepolymer with glycidyl methacrylate or
2-aminoethyl methacrylate. In some embodiments, the aliphatic diol
is HO(CH.sub.2).sub.zOH, wherein z is 2-20.
[0063] In some embodiments, in addition to a curable (e.g.,
photocurable) polymer component, a biomaterial ink comprises one or
more of: a suitable solvent, a photoinitiator, a co-initiator, a
free-radical quencher, a UV-absorber, etc. In some embodiments,
suitable additional components of a biomaterial ink include ethyl
acetate, 1-butanol, Diethyl adipate, 1,6-hexanediol diacrylate,
Diethyl fumarate, Irgacure 819, 2-hydroxy-2-methylpropiophone
(Homp), Camphorquinone, 4-ethyl-N,N-dimethylaminobenzoate, dyes
such as Yellow 5 and Sudan 1, etc. Additional components will be
understood in the field.
[0064] In some embodiments, a biomaterial ink comprises one or more
non-curable polymers or other materials, in addition to the
photo-curable polymer component. In some embodiments, upon curing
of the biomaterial ink, a composite (e.g., noncovalently
association) is formed between the cured polymer component and the
non-curable component. In some embodiments, the non-curable
component is stabilized within the composite by the cured polymer.
Therefore, in some embodiments, biomaterial inks and the cured
composites thereof may comprise curable (or cured) polymer
component and one or more additional compounds, oligomers,
polymers, hydrogels, thermosets etc. For example, biomaterial inks
(and materials formed therefrom) may comprise one or more
biodegradeable polymers to form a composite material. Suitable
biodegradeable polymers include, but are not limited to: collagen,
elastin, hyaluronic acid and derivatives, sodium alginate and
derivatives, chitosan and derivatives gelatin, starch, cellulose
polymers (for example methylcellulose, hydroxypropylcellulose,
hydroxypropylmethylcellulose, carboxymethylcellulose, cellulose
acetate phthalate, cellulose acetate succinate,
hydroxypropylmethylcellulose phthalate), poly(diol citrate) (e.g.,
poly(octanediol citrate), etc.), casein, dextran and derivatives,
polysaccharides, poly(caprolactone), fibrinogen, poly(hydroxyl
acids), poly(L-lactide) poly(D,L lactide),
poly(D,L-lactide-co-glycolide), poly(L-lactide-co-glycolide),
copolymers of lactic acid and glycolic acid, copolymers of
.epsilon.-caprolactone and lactide, copolymers of glycolide and
.epsilon.-caprolactone, copolymers of lactide and
1,4-dioxane-2-one, polymers and copolymers that include one or more
of the residue units of the monomers D-lactide, L-lactide,
D,L-lactide, glycolide, .epsilon.-caprolactone, trimethylene
carbonate, 1,4-dioxane-2-one or 1,5-dioxepan-2-one,
poly(glycolide), poly(hydroxybutyrate), poly(alkylcarbonate) and
poly(orthoesters), polyesters, poly(hydroxyvaleric acid),
polydioxanone, poly(ethylene terephthalate), poly(malic acid),
poly(tartronic acid), polyanhydrides, polyphosphazenes, poly(amino
acids), and copolymers of the above polymers as well as blends and
combinations of the above polymers (See generally, Illum, L.,
Davids, S. S. (eds.) "Polymers in Controlled Drug Delivery" Wright,
Bristol, 1987; Arshady, J. Controlled Release 17:1-22, 1991; Pitt,
Int. J. Phar. 59:173-196, 1990; Holland et al., J. Controlled
Release 4:155-0180, 1986; herein incorporated by reference in their
entireties). Composites of the curable (or cured) polymer and
non-polymeric materials are also within the scope of embodiments
described herein. Such non-polymer components include, but are not
limited to a bioceramic (e.g., hydroxyapatite, tricalcium
phosphate, etc.), nanoparticles (e.g., iron oxide, zinc oxide,
gold, etc.), etc.
[0065] In some embodiments, the curable (or cured) polymer
comprises at least 10% (e.g., 10%, 20%, 30%, 40%, 50%, 60%, 70%,
80%, 90%, 95%, 98%, 99%, 100%) of the biomaterial ink and/or the
resulting cured biomaterial. The aforementioned percentages may be
wt % or molar %.
[0066] In some embodiments, many characteristics of the devices
made with the materials and methods described herein are
customizable. For example, to enable visibility of stents in the
operating room, the radiopacity if the materials was considered. To
enable radiopacity of stents (or other devices and implants), a
large variety of possible materials could be used. In experiments
conducted during development of embodiments herein, visipaque and
Iodixanol have been incorporated devices (e.g., stents). One issue
with stents that are used currently is that a stent exhibiting
radiopacity blocks the view of a variety of scanning techniques
that doctors use to determine the extent of restenosis. Since the
device and stents herein will absorb into the body, restenosis
rates are more easily monitored to determine if and when an
additional follow-up procedure is necessary to protect the
patient's health.
[0067] In additional to radiopacity, in some embodiments, devices
comprise materials to serve as contrast agents. This allows the
devices to be monitored by various biophysical techniques, such as
x-ray, magnetic resonance imaging (MRI), positron emission
tomography (PET), computed tomography (CT), or single-photon
emission computed tomography (SPECT). Any suitable contrast agent
could be incorporated into the materials and devices described
herein. For example, in some embodiments, an iodinated contrast
agent is incorporated into the materials and devices, such as one
selected from the group consisting of iohexyl, iopromide, ioversol,
ioxaglate and iodixanol.
[0068] Other agents may be incorporated into the biomaterial inks,
materials and devices herein. These agents may be covalently
attached to a component of the ink (e.g., the polymer component),
embedded within the material, coated onto a device, etc. Suitable
agents include, but are not limited to: anticoagulants (e.g.,
heparin, Coumadin, protamine, hirudin, etc.), antithrombotic agents
(e.g., clopidogrel, heparin, hirudin, iloprost, etc.), antiplatelet
agents (e.g., aspirin, dipyridamole, etc.), anti-inflammatory
agents (e.g., methylprednisolone, dexamethasone, tranilast, etc.),
anti-proliferative/immunosuppresive agents (e.g., trapidil,
tyrphostin, rapamycin, FK-506, mycophenolic acid), cytostatic drugs
(e.g., paclitaxel, rapamycin, rapamycin analogs (e.g., everolimus,
tacrolimus, etc.), etc.), lipid-lowering agents (e.g., statins),
antioxidants (e.g., probucol, vitamin C, retinoids, resveratrol,
etc.) etc.
[0069] In some embodiments, an mPDC base polymer, and any
polydiolcitrates or methacrylated poly(diol citrates), are
intrinsically antioxidant, which was confirmed by incubating mPDC
(50 mg/mL) in 2,2'-azino-bis(3-ethylbenzothiazoline-6-sulphonic
acid) or ABTS radical solution. mPDC slowly neutralized free
radicals over time with 70% scavenged after 14 days (FIG. 22-Left).
To assess the biocompatibility, UV-cured mPDC films were sterilized
and seeded with vascular smooth muscle cells. Cells could attach
and spread and showed excellent viability after 3 days of cell
culture (FIG. 22-Center). 3D-printed stents (composition: 50% DEF,
47.78% mPDC, 2.2% Irgacure 819, 0.02% Sudan I) degraded over time
upon incubation in PBS at 37 C, with approximately 25% degraded
after 6 months (FIG. 22-Right). These unique characteristics of
mPDC makes it a particularly interest candidate for fabricating the
bioresorbable scaffolds.
[0070] In some embodiments, systems, devices, and methods are
provided for fabricating biomaterial devices (e.g., implants,
stents, etc.) of defined shapes and dimensions from a curable
(e.g., photo-curable) biomaterial ink. Although projection
micro-stereolithography (P.mu.LS) and micro-continuous liquid
interface production (micro-CLIP) are exemplified in Example 1 and
Example 2 below, the scope of embodiments herein are not limited to
such systems and methods. Any suitable systems, devices, and
methods for the controlled application and of biomaterial ink and
conversion of the biomaterial ink into a biomaterial device is
within the scope of embodiments herein. Exemplary systems and
processes, all or a portion of which may be utilized in embodiments
herein, are described in connection with the biomaterial inks and
device-production embodiments herein.
[0071] In some embodiments, systems, methods, and devices from
laser scanning stereolithography techniques are utilized. In such
systems, curing between polymers is induced by
micro-stereolithography, under the action of light. In some
embodiments, a laser scanning unit exposes a defined area on the
surface of the biomaterial ink, in a desired pattern, and in that
way, with a given depth of penetration, hardens a layer of the
pattern to be produced into a solid biomaterial. A displacement
unit in the z-direction provides that the substrate is lowered
layer by layer by the defined layer thickness or the laser focus is
raised. In a processing step, the biomaterial ink, over the
previously produced solid biomaterial layer. This process is
repeated until the desired structure is produced. Laser scanning
stereolithography techniques which may be employed, alone or in
combination with other 3D printing and/or additive manufacturing
systems and processes are further described in Balashanmugan et al.
Procedia Materials Science, Volume 5, 2014, Pages 1466-1472; which
is incorporated by reference in its entirety.
[0072] In some embodiments, systems, methods, and devices from
projection micro-stereolithography techniques are utilized (See,
e.g., Example 1). Projection micro-stereolithography (P.mu.SL)
adapts 3D printing technology for micro-fabrication. Digital micro
display technology provides dynamic stereolithography masks that
work as a virtual photomask. This technique allows for rapid
photopolymerization of an entire layer with a flash of UV
illumination at micro-scale resolution. The mask controls
individual pixel light intensity, allowing control of material
properties of the fabricated structure with desired spatial
distribution. The dynamic mask defines the beam. In some
embodiments, the beam is focused on the surface of a UV-curable
polymer resin through a projection lens that reduces the image to
the desired size. In some embodiments, once a layer is polymerized,
the stage drops the substrate by a predefined layer thickness, and
the dynamic mask displays the image for the next layer on top of
the preceding one. This proceeds iteratively until complete.
P.mu.LS techniques which may be employed, alone or in combination
with other 3D printing and/or additive manufacturing systems and
processes are further described in Zheng et al. Rev Sci Instrum.
2012 December; 83(12):125001; which is incorporated by reference in
its entirety.
[0073] In some embodiments, systems, methods, and devices from
direct inkjet 3D printing techniques are utilized. Direct inkjet
printing systems fabricating a part/device by an additive
manufacturing process. For example, in some embodiments, an ink
delivery system operative to circulate the biomaterial ink, a
printhead associated with the ink delivery system, dispenses the
biomaterial through one or more nozzles based on a defined pattern
(e.g., CAD defined pattern) onto a surface for receiving the
dispensed biomaterial ink one layer at a time. In the case of the
curable (e.g., photo-curable) biomaterial inks herein, the
dispensed ink is exposed to a cure-induced (e.g., light) in order
to produce a layer of solid biomaterial on the receiving surface.
The part/device is formed from a plurality of layers, as the
biomaterial ink is dispensed from the printhead and the ink is
cured in successive layers. P.mu.LS techniques which may be
employed, alone or in combination with other 3D printing and/or
additive manufacturing systems and processes are further described
in Muller et al. Prod. Eng Res. Devel. (2014) 8:25-32; which is
incorporated by reference in its entirety.
[0074] In some embodiments, systems, methods, and devices from
Continuous Liquid Interface Production (CLIP) and/or Micro
Continuous Liquid Interface Production (Micro-CLIP) techniques are
utilized. In CLIP, the continuous process begins with a pool of
photo-curable biomaterial ink. A portion of the pool bottom is
transparent ("window") to light (e.g., UV light). A light beam
shines through the window, illuminating a precise cross-section of
the object. The light converts the biomaterial ink into a sold
biomaterial. The formed object rises slowly enough to allow the ink
flow under and maintain contact with the bottom of the object. In
some embodiments, an oxygen-permeable membrane lies below the ink,
creating a "dead zone" (persistent liquid interface) preventing the
ink from attaching to the window. P.mu.LS techniques which may be
employed, alone or in combination with other 3D printing and/or
additive manufacturing systems and processes are further described
in Dendukuri, D. (2006). Nature Materials 5, 365-369 (2006); which
is incorporated by reference in its entirety.
[0075] The devices, elements, systems, methods, techniques, etc.
from any of the aforementioned 3D printing techniques may be
utilized in any combination in embodiments herein.
[0076] Due to the biodegradable and biocompatible nature of the
materials described herein, the devices and components produced by
the systems, materials, and methods herein find particular utility
in biomedical applications. In some embodiments, devices or
components/parts of devices for implantation into a subject are
produced by the systems and methods described herein. Depending
upon the particular biomaterial selected, the
permanence/impermanence of the particular device may be tailored
(e.g., biodegradation over 1 week, 2 weeks, 1 month, 2 months, 3
months, 4 months, 6 months, 8 months, 1 year, 2 years, 3 years, 5
years, 10 years, or ranges therebetween). Embodiments herein are
not limited by the type of device or implant, or component/part
thereof, that is produced by the systems, materials, and methods
described herein. Exemplary implants, devices, etc. that may be
manufactured by the systems, materials, and methods described
herein, or may have a part or components that may be manufactured
by the systems, materials, and methods described herein, include,
but are not limited to: stents, stent-grafts, grafts, vascular
grafts, shunts, screws, nails, threads, clasps, tubes, catheters,
patches, plates, sheets, meshes, ports, rings, prostheses, contact
lenses, ocular implants, cardiovascular implants, pacemakers,
orthopedic implants, sockets and counterparts, etc.
[0077] Although the systems, materials, and methods described
herein have particular utility in biomedical applications,
embodiments within the scope herein are not so limited. In some
embodiments, devices and parts produced by the systems, materials,
and methods herein find use in the fields of veterinary medicine,
laboratory research, microfluidics, environmental science,
industrial, and other applications.
[0078] In some embodiments, materials and devices are provided with
self-expanding material properties, shown in FIG. 20. In some
embodiments, the devices (e.g., stents) are designed and printed in
the fully expanded state and then sheathed or compressed into a
collapsed conformation (e.g., within a catheter). In some
embodiments, the self-expanding character is useful for implantable
devices (e.g., stents), for example, for derives that are implanted
into peripheral arteries, especially when those arteries are near
areas of the body that can be collapsed by external forces such as
the arteries within the thigh and near the knees.
[0079] In some embodiments, devices are provided comprising
balloon-expandable designs. In some embodiments, balloon-expandable
devices comprise a PLLA-based material that is plastically
deformable. Balloon-expandable stents, for example, are used almost
exclusively in cardio implants, and are preferred in that field due
to their improved flexibility and stronger radial strength.
EXPERIMENTAL
[0080] While the compositions and methods described herein may find
use with any suitable photo-polymer based additive manufacturing
devices/techniques (e.g., Laser Scanning Stereolithography,
Projection Stereolithography, Ink-Jet Printing, Continuous Liquid
Interface Production (CLIP), etc.), two printer types are
exemplified below to demonstrate successful fabrication, imaging
and mechanical testing of the designs to validate them. Exemplary
methodologies and utilities of each printing type are explained in
the examples below.
[0081] The following examples provide exemplary embodiments within
the scope herein, and data representing experiments conducted
during development of such embodiments. These data exemplify the
embodiments herein, but should not be viewed as limiting on the
scope of embodiments herein.
Example 1
Projection Micro-Stereolithography
[0082] A biodegradable biomaterial ink was formulated with
biodegradable methacrylated poly(diol citrate)s and enables rapid
fabrication of endovascular devices (e.g., stents) via projection
microstereolithography technique. As exemplary devices, stents with
various microstructures were printed in a resolution of 20 um using
CAD modeling. mPDC stent showed a compliant compressive strength
and flexibility. Compared with bare metal stent, numerical
simulation showed the experimental results differed by
approximately a factor of 5, the 350 um stent best approximates the
Nitinol BMS stent.
Polymer Synthesis and Characterization
[0083] The synthesis of methacrylated poly(1,12-dodecanediol
citrate) is depicted in FIG. 22. A similar scheme is applicable to
the methacrylation of other poly(diol citrates), which find use in
other embodiments herein.
[0084] Citric acid (76.8 g; Sigma) and 1,12-dodecanediol (40.4 g;
Sigma) were added to a flask and heated to 165.degree. C. under
nitrogen atmosphere. After melting, the reaction was continued for
an additional 30 minutes at 140.degree. C. The viscous
poly(1,12-dodecanediol citrate) (PDC) pre-polymer is dissolved in
100-150 ml ethanol and purified by precipitation in 1000 mL of
deionized water (Millipore water purification system), then
freeze-dried for at least 72 hours. Subsequently, 22 g PDC was
added to 180 mL tetrahydrofuran (Sigma) for dissolution, then 816
mg of imidazole (Sigma) and 17.04 g of glycidyl methacrylate
(Sigma) was added and heated to 60.degree. C. for 6 hours then
placed on a rotary evaporator for 30 minutes at 60.degree. C. After
methacrylation, mPDC was purified using 900 mL of deionized water
twice, then centrifuged in 50 mL vials for 5 minutes at 3500 rpm
followed by freeze drying for 24 hours. The purified mPDC polymer
was characterized using a Bruker Ag500 NMR spectrometer at ambient
temperature, using DMSO-d6 as solvent, and tetramethylsilane (TMS)
as the internal reference.
Biomaterial Ink Formulation and Rheological Characterization
[0085] The viscous mPDC polymer was diluted with different
chemicals such as ethyl acetate (Anhydrous, 99.8%; Sigma),
1-butanol (ACS reagent, >99.4%; Sigma), Diethyl adipate
(ReagentPlus.RTM., 99%; Aldrich), 1,6-hexanediol diacrylate
(Technical grade, 80%; Aldrich) and Diethyl fumarate (98%;
Aldrich), 0.1-5wt % amounts of initiators such as Irgacure 819,
2-hydroxy-2-methylpropiophone (Homp) and Camphorquinone were
formulated into mPDC solution for curing at different wavelengths.
Compatibly, 4-ethyl-N,N-dimethylaminobenzoate was used as a
co-initiator to accelerate the reaction, a number of dyes such as
Yellow 5 and Sudan 1 served as a free radical quencher or UV
absorber. The UV/Vis absorption spectra of different initiators
were recorded in an Aligent Cary 100 spectrophotometer. Rheological
measurement of mPDC solution in ethyl acetate was carried out on a
TA instruments DHR rheometer with a 20 mm 4.degree. cone peltier
plate geometry and solvent trap cover to minimize sample
evaporation. A flow ramp experiment was performed for 0.1 to
142.665 rad/s at 25.degree. C. and 37.degree. C. to determine the
dynamic viscosities of pure mPDC and mPDC solution with 5 wt %, 10
wt % and 15 wt % ethyl acetate. Viscosity changes as a function of
shear rate were assessed via rheometry.
Projection Microstereolithography Printer Design and
Fabrication
[0086] Projection microstereolithography (P.mu.SL) builds
microstructures from a photo-curable biomaterial ink in a
layer-by-layer fashion directly from a 3D CAD design. Each layer is
cured in a single exposure by using a liquid crystal display (LCD)
panel as a dynamic mask for the UV light. This allows for a drastic
reduction in fabrication time compared with conventional 3D
printing process, which fabricates 3D structures in a
point-by-point scanning fashion.
[0087] An exemplary process flow is depicted in FIG. 7. Prior to
fabrication, a photo-curable biomaterial ink was formulated as
described in the section below. The CAD structure is sliced into a
series bitmap images using a MATLAB code developed specifically for
this system. The UV absorber and light intensity concentration is
tuned to obtain a curing depth of 20 microns, determining the
necessary slicing layer thickness.
[0088] The silicon wafer is then aligned with the top of the
biomaterial ink layer, and the 160 liter P.mu.SL chamber is filled
with nitrogen gas. This reduces the concentration of oxygen within
the chamber and ensures optimal solidification and resolution of
the photo-curable biomaterial ink. Afterwards, the layer building
process begins. The first sliced bitmap image is displayed on the
dynamic mask (in this case, a 1400.times.1050 pixel array), and the
wafer drops by 20 microns. The system then waits for 30 seconds for
the biomaterial ink to settle. The UV lamp is turned on for 20
seconds, reflects off a beam splitting mirror, passes through a
reduction lens and finally projects onto the surface of the
biomaterial ink in high resolution, with each pixel corresponding
to 7.1.times.7.1 .mu.m.sup.2 repeats for each bitmap layer in the
fabrication. The micro-structure is then removed from the P.mu.SL
machine, cleaned with isopropyl alcohol (IPA), dried under a low
flow rate nitrogen gun. At this point, the biomaterial ink within
the structure has not completely solidified. To finish the curing
process and bring the biomaterial ink to its final state, the
structure is further exposed to UV for post-curing.
Stent Design
[0089] Stent design with various microstructures were prepared
using the SOLIDWORKS CAD software (Waltham, Mass.). Sinusoidal
formed wire, helix wrap and meshed tube was created and printed
along the circumference layer-by-layer with length.times.outer
diameter.times.thickness. Various parameters such as 300 um, 350
um, 400 um and 500 um in thickness or 9.0 mm, 16 mm and 21 mm in
length were investigated. Typically, a stent pattern was chosen to
be a triangular truss structure along the circumference with each
new row connected via vertical supporting rods, as shown in FIG.
3a. Each new row was shifted to allow the lowest point of the upper
row to be in line with the highest point of the bottom row. These
points were then connected by vertical beams that gives the
appearance of hexagonal holes across the face of the cylinder. To
avoid misalignment and a floating point at the low point of the top
row, vertical support rods were placed at low and high point
section for fabrication.
[0090] The rods with smaller cross section act as removable support
structure that were removed after fabrication was completed, outer
diameter of stent was given a set value of 5.20 mm. Stent strut
thickness was set to 350 um, the individual "true support" stent
rod diameter was also set at 350 um and a height of 550 um tall.
The "removable support" material rods were set to a value of 100 um
with a height of 300 um. Further support rods of 150 um diameter
and 300 um tall were placed at the bottom of the stent to allow
easy removal from the base. The entire stent was built on a square
base of 5.5 mm.times.5.5 mm by 500 um tall. This overall design was
initially chosen in order to verify the capability of the P.mu.SL
system to manufacture such structures as stents. Optimization to
this design and other design changes was performed.
Morphological Assessment of the Stents
[0091] Samples of printed stents were observed in high vacuum mode
(<10-4Torr) with 10 kV operation voltage by utilizing FEI Quanta
environmental scanning electron microscopy (ESEM) without polish
and coating.
Mechanical Testing
[0092] Mechanical compressive tests of mPDC stents were conducted
according to ASTM D2412-11 by parallel-plate loading on an Instron
5544 mechanical tester equipped with 500 N load cell at a rate of
100 mm/min (Instron, Canton, Mass.). Radial compression testing was
performed by compressing mPDC stents a total of 2 mm corresponding
to 33% to 50% displacement depending on stent outer diameter. A
three-point bend test apparatus (a cylindrical actuator in the
middle of two cylindrical end-supports at a distance of 20 mm) was
used for flexibility testing, which was performed according to ASTM
F2606-08 on a MTS Sintech 20/G Universal Testing Machine with 210 N
load cell at a crosshead rate of 10 mm/min (Sinotech, Portland,
Oreg.). The maximum bending angle was set at 48.degree..
Numerical Simulation
[0093] Numerical simulation for three-point bending and
parallel-plate compression of stents were performed utilizing the
SOLIDWORKS (Waltham, Mass.) and ANSYS workbench (Cecil Township,
Pa.) softwares. Three thicknesses of 300 um, 350 um, and 400 um and
length of 21 mm of stents were examined for both the parallel-plate
compression and three-point bending simulations. To simulate
3-point bending, the stent was fixed on one side and two regions
near the edges while forces were added along half the length of the
stent, the displacement field was analyzed. In the parallel-plate
compression, the forces were applied in a slim region along the
length of stent.
Polymer Synthesis and Characterization
[0094] Citric acid is a multifunctional monomer in the Kreb's cycle
that is easily reacted with various diols to form a crosslink
elastomer in the absence of exogenous catalysts (ref. 26;
incorporated by reference in its entirety). Under a controllable
condition and procedure, the synthesized PDDC prepolymer was
uncrosslinked and was dissolvable in several solvents such as
ethanol, acetone, dioxane, etc (ref. 28; incorporated by reference
in its entirety). In basic conditions, glycidyl methacrylate was
used in an epoxide ring-opening reaction to attack the unreacted
carboxylic groups of citric acid using imidazole as a catalyst.
Methacrylate was successfully introduced to the PDDC backbone. A
novel mPDC polymer was obtained as determined by .sup.1H NMR
spectrum with evidence of proton peaks for citrate residues (1) and
methacrylate residues (5 and 6) (FIG. 1). The multiple peaks at
2.79 ppm were assigned to the protons in --CH2-- from citric acid,
and the peak at 1.84 ppm was assigned to --CH3 in methacrylate
unit. The molar composition of mPDC calculated from the signal
intensities of both protons was approximately 1:1 of citric
acid/methacrylate. mPDC polymer immediately forms a solid by
photopolymerization after mixing with a photoinitiator as shown in
FIG. 1.
Biomaterial Ink Formulation and Rheological Characterization
[0095] mPDC polymer is easily dispersed and formulated in different
chemicals such as ethanol, acetone, dioxane, ethyl acetate,
1-butanol, Diethyl adipate, 1,6-hexanediol diacrylate and Diethyl
fumarate, etc. mPDC viscosities do not change significantly in a
shear rate from 1 to 150 l/s, at 15.5.+-.0.4 Pas as shown in FIG.
2c. Upon adding different amounts of ethyl acetate, the mPDC
solution remains flowing stable, the viscosities remarkably
decrease over shear rate with the increasing ethyl acetate, from
8.0.+-.0.5 Pas in 5 wt % to 1.50.+-.0.04 Pas in 15 wt %. However,
all the viscosities of the polymer and solution decrease over
temperature, heating can increase the flowability of both polymer
and solution.
[0096] It was observed in the experiments conducted during
development of embodiments herein that all the initiators,
co-initiators, and free radical quenchers are easily dissolved in
the mPDC solution forming a homogenous solution and quickly forming
a solid upon exposure to light. FIG. 2a showed the UV/Vis
absorption of different initiators such as Irgacure 819,
Camphorqinone and 2-hydroxy-2-methylpropiophone separately in 370
nm, 470 nm and 340 nm, with the concentration dependence. After
being cured with Camphorqinone at 470 nm, the mPDC stent in 0.5 mm
thickness showed complete compliance with BMS in compressive
strength in FIG. 2d, compressive modulus of mPDC stent in 0.75 mm
increased to 10.64.+-.3.6 MPa. Similarly, to increase the
resolution of projection microstereolithography printing, 2.2 wt %
Irgacure 819 was used as the photoinitiator and 0.18-0.22 wt %
Sudan 1 as the UV absorber after a series of optimization. Irgacure
819 can easily bind both HDDA and mPDC independently at a molecular
level, while Sudan I absorbs UV light at 405nm provided by the
printer to control the curing depth.
Projection Microstereolithography Printer Design and
Fabrication
[0097] Projection microstereolithography printer design was based
on digital micromirror device (DMD, Texas Instrument) as a dynamic
mask at 1400.times.1050 pixels that is the core of this technique
to use a spatial light modulator. The modulated light was
transferred through a reduction lens (CoastalOpt 60 mm UV-VIS-NIR
lens, JENOPTIK Optical System Inc) to the surface of biomaterial
ink with the reduced feature sizes, each pixel in the dynamic mask
is focused down from original dimensions (object size) of 10
um.times.10 um to an image size of 7.1 um.times.7.1 um, the
magnification is approximately 1.4. The biomaterial ink can be
cured at a 2D pattern in a single exposure and stacked in a series
of closely spaced horizontal planes programmed by a 3D CAD model.
In the projection microstereolithography printer, the intensity of
UV light is controlled by the current input into the system with
0.4 A at 405 nm, the measured intensity is 0.03 mW. Typically, the
curing time for HDDA stent is 12 seconds per layer and 20 seconds
per layer for mPDC stent. With this bottom-to-top fabrication, the
biomaterial ink enables printing the stents with high resolution of
7 um pixel in a curing depth of 20 um. The cured biomaterial ink
has strong enough mechanical properties to enable 350-400 um struts
over a 21 mm stent design height, as shown in FIG. 3c-d and FIG. 4,
each layer is 20 um in depth with precise edges.
Morphological Assessment of the Stents
[0098] In this process of bottom-to-top microfabrication, various
microstructures in the stents were also showed in FIG. 3c-d and
FIG. 4. Sinusoidal wire and fiber mesh were stacked in circular and
rectangular layers with 20 um height. In FIG. 3, SEM images showed
sinusoidal stent was interconnected with bridges in 0.55 mm as
designed as vertical support rods. Experiments conducted during
development of embodiments herein demonstrate that projection
microstereolithography can print the stents with various
microstructures.
Mechanical Testing
[0099] Parallel-plate compression and 3-point bending experiments
were performed to determine the mechanical properties of stents.
Unlike in-situ mPDC tubes, the mechanical properties of printed
stent are significantly affected by stent design and its
microstructure. mPDC significantly change HDDA compression strength
and make it more flexible to match Nitinol BMS, as shown in FIG.
6b. By compressing 2 mm from 5.5 mm in outer diameter, no complete
rupture was found so that this closed microstructure resists the
mechanical fracture of stent.
Numerical Simulation of the Stent Design
[0100] To accomplish the simulation for three-point bending, on one
half of the stent circumference, two regions near the edges of the
stent were fixed, while forces were added along half the length of
the stent on the opposite side of the fixed area. A typical
displacement field from a 400 um thick stent is presented in FIG.
5d. From an applied force of 0.5N onto the 400 um stent, the
resulting maximum displacement of 7.945 mm was observed where
forces were applied. The displacement on the opposite end of the
stent was 3.783 mm. For all three stent thicknesses, the maximum
displacement was plotted against the force applied to the stent
(FIG. 5c). For increasing stent thickness, the necessary applied
force to displace the stent increases. The primary properties
analyzed in these simulations were the range of forces that these
stents are predicted to be usable. The stent was considered
"unusable" when the point of maximum displacement is within 1 mm
from the point on the opposite side of the circumference of the
stent. This was determined from the following equation:
(Dd+d2)-d1;
where Dd is the outside diameter or the stent, d2 is the
displacement of the point opposite the point of maximum
displacement, d1 is the maximum displacement value. With increasing
thickness (.mu.m), the Applied Force necessary to make the stent
unusable increased (FIG. 5d). These Applied Force values for the
300 um, 350 um, and 400 um thick stents were 0.18N, 0.325N, and
0.555N, respectively.
[0101] The parallel-plate compression was simulated on the three
stent designs. With increasing force, there is nonlinear contact
between the plates and the new deformed surface of cylinder. The
parallel-plate compression analysis was done for this study by
fixing a slim region along the length of the cylinder be fixed.
Particular faces on the cylinder's opposite side were subjected to
equal forces. As with three-point bending, the range of forces that
the stents were "usable" were analyzed (FIG. 6a) and the Maximum
Usable Applied Force was plotted in relation to stent wall
thickness (FIG. 6b). With increasing stent thickness, the necessary
force to cause deformation of the stent walls increased. The
numerical results and the experimental results differed by
approximately a factor of 5. A scaling factor of 5 was used in
order to compare the experimental results and the numerical
results. FIGS. 6a and 6c represent Applied Forces multiplied by the
scaling factor.
[0102] With the scaling, the numerical simulation becomes more
directly comparable to the experimental tests. The "HDDA" stent
used for the experimental tests was a design that was 500 um in
thickness for both the stent walls and all supporting rods. This
slightly differs from the design that is numerically evaluated. The
design that is being evaluated numerically has some support
structures of (100 um). At 2 mm compression, the HDDA stent needed
approximately 4N of applied force to cause displacement. At 2 mm of
compression, the 500 um design needed approximately 3.5N of force.
This difference could be attributed to the inclusion of smaller 100
um support rods. The 350 um stent best approximates the Nitinol BMS
stent. In both cases, approximately 1N of force is necessary to
compress the structure 2 mm.
Example 2
Micro-CLIP Additive Manufacturing Process
Design, Print Custom Design, Tailor Performance
[0103] Using the exemplary manufacturing process described below,
7.1 um lateral resolution was obtained. The combination of
manufacturing process, material, and design flexibility allow for
the custom fabrication of stents to fit the needs of a particular
subject or application.
[0104] Micro-CLIP manufacturing method is based on a similar
methodology as Projection Micro-Stereolithography. In some
embodiments, the Micro-CLIP system is capable of printing up to 200
times faster than projection stereolithography method.
[0105] With a single Micro-CLIP printer devices were generated at
the necessary scale for low-volume manufacturing. When using
projection stereolithography (PuSL), 16 hours of time were required
for a single print. With Micro-CLIP a new 20 mm length stent can be
printed in just five minutes. By Micro-CLIP, the slowest prints
tested took only seventy minutes, which is nearly 15 times faster
than the PuSL system for a high resolution object of 1000 layers.
This time is further reducible through the use of properly
optimized material, light source, and dead zone. This technology
has additional advantages including the ability to work with a
broader array of polymer materials and each print has isotropic
material properties. However, compared to P.mu.SL, Micro-CLIP has
weaker provided mechanical properties under compression. With
P.mu.SL, the maximum stent length possible was 20 mm. With
Micro-CLIP, the maximum stent length achieved is 48 mm, with
significantly greater length achievable. Devices much taller than
200 mm are achievable with this technology depending upon the
materials used and the structure to be printed.
[0106] Micro-CLIP additive manufacturing provides for the
fabrication of microstructures from a photo-curable biomaterial ink
in a layer-by-layer fashion directly from a 3D CAD design. Each
layer is cured in a single exposure using a digital
micromirror-device (DMD) as the dynamic mask for the UV light. This
differs from the P.mu.SL system which uses a liquid crystal
display. The liquid crystal display is not able to withstand the
high power UV required for the Micro-CLIP process. This allows for
a dramatic reduction in fabrication time compared with conventional
3D printing processes, which fabricate 3D structures in a
point-by-point scanning fashion. In addition to fabrication of an
entire surface area at once, CLIP operates under nearly continuous
motion. In a P.mu.SL process, after the first layer forms, the
fabricated part is typically dipped back into the liquid resin bath
and then raised so that only a single .about.5-20 um layer of
liquid is on top of the part, then time is allowed for the material
to settle, a process that can take 30 seconds to two minutes per
layer depending on the material viscosity. That entire process is
eliminated in CLIP. With CLIP, the platform moves upwards at a
nearly constant speed, only stopping for 10 ms-100 ms between each
layer, dramatically reducing part print speed. Additionally, a
higher intensity of UV light is used which enables photocuring each
layer of the part in dramatically less time.
CLIP Process Flow
[0107] Prior to device (e.g., stent) fabrication, a photo-curable
biomaterial ink was formulated as described in the section below.
The CAD structure is sliced into a series of bitmap images using a
MATLAB code developed specifically for this system. The UV absorber
and light intensity concentration were tuned to obtain a curing
depth of 20 microns to tune the finalized surface finish of the
part. During the alignment process the Teflon AF2400 thin film was
aligned to be placed 20 um below the focal plane of UV intensity.
During the print process the build platform then drops down until
it comes in contact with the Teflon AF2400 thin film, contact is
determined via a force sensor built into the platform. The purpose
of the Teflon AF2400 thin film is to control the oxygen flow rate
that makes contact with the liquid resin. Oxygen inhibits the
photo-polymerization reaction and by allowing just a small amount
into the bath a dead-zone forms. The printing process then begins
and the first sliced image is displayed on the digital micromirror
devices (in this case a 1980.times.1050 pixel array). The system
begins moving upwards at the desired user controlled speed (80 um/s
for example) until the system has moved upwards 20 um. The system
then briefly stops, switches images to the second sliced image,
waits for 10-100 ms to ensure a full switch of the image, and then
begins moving again at the user controlled speed. Speed, UV
intensity, and image are dynamically controllable and modulatable
at each individual layer of the print. Layer thickness does not
have to be 20 um, it can be as low as 100 nm. This process
continues, with the platform continuing to move up and new images
continuing to be displayed until the entire part is completed.
[0108] In terms of the light path when the UV LED is turned on, the
light first passes through a collimating lens, through a light gate
and then reflects off a digital micro-mirror device which contains
millions of tiny mirrors. The reflected light passes back through
the light gate, through a focusing lens and beam-splitter and off a
90 degree mirror before ending at the focal plane with each pixel
corresponding to 7.2.times.7.2 um 2.
[0109] After the part was complete, the micro-structure was then
removed from the machine, excess material was cleaned off with a
chem-wipe and the part was left in a dionized water bath for a few
hours to remove any excess material. To improve mechanical strength
the parts were then removed from the water bath, dried under a
nitrogen gun and post-cured under an intensity of 350 mW/cm.sup.2
for 6 minutes (3 minutes on each side).
Resolution Accuracy
[0110] Resolution of the fabrication systems is affected by several
variables from both the fabrication system and the fabrication
material. Potential variables include the following: speed of
fabrication, light intensity, amount of pause at each fabrication
layer (exposure time), concentration of UV absorber in material,
and concentration of photoinitiator in the material. Several
fabrication tests were performed that varied several of the
parameters listed above. Shown in FIG. 10 are the dimensional
differential vs. light intensity plots from four tests that were
performed. For these tests, the fabricated dimensions of the stents
were compared against the intended stent design dimensions. "Base"
stent design has an intended dimension of 151.4 um strut thickness
in the axial and lateral (planar) directions. Dimensional
differential is the percentage difference between the actual
fabricated dimension and the intended dimension. Values below the
X-axis represent the fabricated dimension is a certain percentage
smaller than intended (underexposure) and values above the X-axis
representing the fabricated dimension being a certain percentage
larger than intended (overexposure). The X-axis represents fully
accurate dimension resolution (correct exposure). Light intensity
was measured as the percentage of the system's maximum intensity.
Photoinitiator and UV absorber used in these tests were Irgacure
819 and Sudan 1, respectively. Exposure time (pause of machine at
each fabrication layer) was either 1 ms or 10 ms. Fabrication speed
was fixed at 5 um/s. Fabricated dimensions were acquired from
scanning electron microscopy and imageJ software and represent an
average along the length of the stent. From these tests areas where
accurate resolution could be achieved were identified for each
material.
[0111] FIG. 9(a) represents a test with a biomaterial ink resin
containing 2.2% photoinitiator, and 0.02% absorber concentrations.
The exposure time for this set of stents was 1 ms. From this test
it was observed that the correct axial exposure was achieved at
approximately 13% or 14% intensity. Correct lateral exposure was
achieved at 17-18% light intensity. FIG. 9(b) represents a test
with the absorber concentration increased to 0.04% and exposure
time increased to 10 ms. At 10% light intensity, axial differential
was only approximately 5%, a decrease to 8% or 9% light intensity
could potentially be give dimensional accuracy. Correct lateral
exposure appears to be achievable at 15% light intensity. In the
test represented by FIG. 9(c), the absorber was increased to 0.08%
concentration, with other variables unchanged. Lateral and axial
differentials are very similar in this test. Both axial and lateral
dimensional accuracy appears to be best at approximately 17-18%. In
the last test, FIG. 9(d), initiator was reduced to 1% concentration
and a polymerization accelerant was added [ethyl
4-(dimethylamino)benzoate] at 1% concentration. Lateral correct
exposure was achievable at approximately 19% or 22-23% intensity.
Correct axial exposure was achieved at approximately 26-27%
intensity. In this last case, the stents appear to laterally
overexpose at lower light intensity than in the axial direction.
For cases where there is overexposure in the axial direction at
lower intensity than the lateral direction, more pixels could be
added to each projected cross section to make all directions
dimensionally accurate. In the case where the lateral direction
experiences overexposure at lower intensity than the axial
direction, a reduction in projected cross section pixels
compensates for lateral overexposure.
Material Flexibility:
[0112] A broad variety of liquid polymer materials function well
within these additive manufacturing processes. Solvents including
Ethanol and Ethyl Acetate have been used to replace Diethyl
Fumarate in the material composition of each individual stent.
Because Ethanol has a lower viscosity than Diethyl fumarate, less
Ethanol is necessary within the final material to match the
viscosity requirements for printing. Ethanol and/or Ethyl Acetate
improve the biocompatibility of the process. In addition to
changing the solvent used, the UV Absorber, Sudan I can be changed
to Benzotriazol, a UV absorber that is nearly transparent in the
visible spectrum and causes the printed object to look clear to the
human eye. A large variety of photo-initiators are compatible with
this process including but not limited to Irgacure 819, Irgacure
651, Irgacure 369, Irgacure 184, Irgacure 2959, Irgacure 1173,
2-hydroxy-2-methylpropiophone (Homp) and Camphorquinone.
Transparent materials are being used to create a look of
cleanliness for both the surgeon and the patient and improve the
aesthetic quality of the device.
Design Flexibility
[0113] AM processes allow for excellent design flexibility and
tunability. With both P.mu.SL and CLIP processes a base stent
design can be experimentally tested and quick design iterations are
possible. The ability to free form fabricate structures with very
high resolution within the span of at most a few hours (P.mu.SL) to
as low as a few minutes (CLIP) allows for very fast direct
experimental testing and design iteration. Depending on the
patient's needs, such as size and location of lesion within the
vessel or vessel geometry, these manufacturing processes
accommodate changes in a base stent design to a complete custom
design. If large radial strength is needed, wall and strut
thicknesses are editable. If more flexibility is needed, strut
connector design is edited. Instead of having a standard set of
sizes for vascular diameter, AM processes allow for a specially
made stent to fit the particular vessel. While AM processes have
certain advantages in terms of flexibility compared to other
manufacturing processes, AM processes still have their own
requirements. For stereolithography based manufacturing (scanning,
projection, CLIP), each fabrication layer must be connected to a
previous fabrication layer, a support fabrication post, or the
build platform. If a design does not account for this requirement,
the printed structure will have structural defects. To accommodate
this requirement, current stent designs have the low point of each
strut ring connected to some portion of the connector ring below
it. Two designs that have been created and parallel-plate
compression tested. Our "Base" design shown in FIG. 11 below and
Arrowhead design in FIG. 12 have been compression tested, while the
Flexibility Optimized Base design (FIG. 14) is a conceptual design
that has not been mechanically tested yet. The Base design was
created to be closely packed to increase radial strength, while the
"S" shaped connections were added to provide reasonable
flexibility. For validation of mechanical properties, a base design
was made similar to a design on the market. Designs made for
patients can be tailored to suit the patient's needs. A unit length
of the "Base" stent design consists of 12 unit cell elements across
the circumference of the stent. This unit length of the stent could
then be added to one another until the desired full length was
obtained. Strut thickness of the Base design was set to be 151.4
um. The angle between struts was set to be 60 degrees. Lastly, the
stent wall was given a thickness of 500 um for the bulk of
mechanical testing.
[0114] For greater emphasis on radial strength, another compact
design was created with simplified connections between struts. FIG.
12 below shows the CAD drawings as well as SEM micrograph of the
Arrowhead stent design. This design also has a 60 degree angle
between struts. Connector thickness and smallest tested strut
thickness was 150 um. Unit length of the stent consisted of 8 unit
cells. Typical unit cell of this design is shown in FIG. 12(c). As
with the base design, unit lengths of the Arrowhead design could be
attached to one another in CAD software until desired length is
obtained.
[0115] For optimization, metamodels are created of how each design
parameter affects the objective and constraint functions.
Constraint functions may be failure stresses, patient vessel
geometric constraints, and fabrication constraints. Metamodels may
be created from data collected via FEM modeling or experimental
data. Following is an example of parametric optimization performed
on a stent design made during experiments conducted during
development of embodiments herein is shown. A parametric
flexibility optimization was performed on the Base design template
to make a stent for more diverse applications. The previously
described stents were mainly designed to favor strength rather than
flexibility. Flexibility is a key component of stents, as
vasculature may curve suddenly, and the stent needs to be able to
be potentially inserted in variety of geometric areas. The design
variables that were varied for study were the strut angle (.THETA.)
and the Connector Height (H). The connector thickness (t) was given
a fixed relationship with the Connector Height, with t being 20% of
the height (FIG. 14(a)). Stress analysis was performed using ANSYS
FEA software. The objective function that was to be optimized was
known as the Flexibility Metric (FM), which was defined as the
integral of Moment vs. Curvature index graph (Pant, S.; Bressloff,
N W; Limbert, G. Biomech. Model Mechanobiol. (2012) 11;
incorporated by reference in its entirety). FM represents a value
to be minimized as it implies that for a particular curvature index
a smaller applied moment is required. Design of experiments was
obtained via Latin Hypercubes in iSight optimization software,
which gave 20 design points of interest. Nineteen of these points
were created in CAD. The design space of (.THETA.) was chosen to be
40 and 110 degrees and the design space of "H" was chosen to be
between 250 um and 1 mm.
[0116] In ANSYS, a unit length of the stent was subjected to
opposing moments (in the out of plane axis) at both axial ends. The
moment was varied between 0 to 0.15N*mm. The stent was constrained,
such that, one end face was completely fixed, while the opposite
end was allowed to deform in the axial direction. The failure
criterion was given a failure criterion of 10 MPa. As moments were
applied, the von Mises stress and resulting deformation readings
were collected. The deformation angle was obtained via the arcsine
of the axial deformation divided by the radius of the stent. This
value allowed the calculation of the Curvature Index (CI) which was
defined as .PHI. divided by the stent unit length (L.sub.unit)
(FIG. 13(b). A metamodel was created using iSight software and
contour plot of the FM was made, which is shown above in FIG.
13(c). The gray shaded region to the left of the graph showed where
failure by exceeding the 10 MPa was likely. The dark shaded
elliptical region represents an area where a minimum of FM could be
found. To find this minimum, adaptive simulated annealing and the
multi-island genetic algorithms were utilized. Both algorithms
appeared to converge to a minimum FM value (FM=4.305E-4), which
corresponded to design inputs H=0.825 mm and .THETA.=68.83 degrees.
This optimized design is shown in FIG. 14 below. This design was
successfully fabricated. In addition to the values of connector
height (H) and strut angle (.THETA.), this design differed from the
"base" design by including thinner wall (400 um), slightly larger
strut (200 um) and reduction in circumferential elements (8, rather
than 12).
Mechanical Properties
[0117] Mechanical tests were performed by mean of radial
compression to 25% of the stents' initial outer diameter, using an
Instron 5544 mechanical tester according to ISO 25539.
[0118] In FIG. 15, it can be seen that the radial strength of stent
can be increased by increasing the UV intensity used during
printing. The radial strength decreases with an increase in Sudan I
concentration, giving flexibility in strength by changing the UV
absorber content.
[0119] From the FIG. 16, it can be seen that post-curing more than
doubles the mechanical strength of the 3D-printed biomaterial ink
stents, but that the post-curing time barely affects the mechanical
properties. Thus, the majority of flexibility in mechanical
properties is possible at the printing stage.
[0120] As can be seen in FIG. 17, there is a strong correlation
between the dimensional and mechanical properties of stents,
indicating that a larger material footprint increases radial
strength of stents. However, the relative content of UV absorber
affects the nature of the correlation. Generally, higher UV
absorber concentrations lead to a faster increase in mechanical
strength with increasing dimensions.
[0121] FIG. 18 demonstrates that for the Arrowhead design the
radial strength does not depend on the strut dimensions, but is
strongly dependent on the wall thickness.
[0122] FIG. 19 indicates that biomaterial ink stents may be printed
using an accelerator compound like EDAB. EDAB accelerates the rate
of radical formation for polymerization initiation.
REFERENCES
[0123] The following references, some of which are cited above by
number, are herein incorporated by reference in their entireties.
[0124] 1. Gundogan, B., et al., Bioabsorbable Stent Quo Vadis: A
Case for Nano-Theranostics. Theranostics, 2014. 4(5): p. 514-533.
[0125] 2. Ulrich Sigwart, M.D., Jacques Puel, M.D., Velimir
Mirkovitch, M.D., Francis Joffre, M.D., and Lukas Kappenberger,
M.D., Intravascular Stents to Prevent Occlusion and Re-Stenosis
after Transluminal Angioplasty. New England Journal of Medicine,
1987. 316: p. 6. [0126] 3. Serruys P W, K. M., Ong A T.,
Coronary-artery stents. New England Journal of Medicine, 2000. 354:
p. 13. [0127] 4. Hermawan, H., D. Dube, and D. Mantovani,
Developments in metallic biodegradable stents. Acta Biomater, 2010.
6(5): p. 1693-7. [0128] 5. Regar E, S. G., Serruys P W., Stent
development and local drug delivery. British Medical Bulletin,
2001. 59: p. 21. [0129] 6. Ranade S V, M. K., Richard R E, Chan A
K, Allen M J, Hel-mus M N., Physical characterization of controlled
release of paclitaxel from the TAXUS Express drug-eluting stent. J
Biomed Mater Res, 2004. 71A: p. 10. [0130] 7. Schmitz K P, G. N.,
Lobler M, Behrend D, Schmidt W, Sternberg K., Drug-eluting stent
technologies for vascular regeneration. International Journal of
Materials research, 2007. 98: p. 6. [0131] 8. Kedia G, L. M., Stent
thrombosis with drug-eluting stents: a re-examination of the
evidence. Catheter cardiovascular intervention, 2007. 69: p. 8.
[0132] 9. Garg S, S. P., Coronary stents: looking forward. Journal
of the American College of Cardiology, 2010. 56: p. 34. [0133] 10.
Mei-Chin Chen, H.-W. T., Yen Chang, Wei-Yun Lai, Fwu-Long Mi,
Chin-Tang Liu, Hen-Sheng Wong, and Hsing-Wen Sung, Rapidly self
expandable polymeric stents with shape memory property.
Biomacromolecules, 2007. 8: p. 7. [0134] 11. Grabow, N., et al., A
biodegradable slotted tube stent based on poly(L-lactide) and
poly(4-hydroxybutyrate) for rapid balloon-expansion. Ann Biomed
Eng, 2007. 35(12): p. 2031-8. [0135] 12. Shih-Jung Liu, F.-J. C.,
Chao-Ying Hsiao, Yi-Chuan Kau, Kuo-Sheng Liu, Fabrication of
Balloon-Expandable Self-Lock Drug-Eluting Polycaprolactone Stents
Using Micro-Injection Molding and Spray Coating Techniques. Annal
of Biomedical Engineering, 2010. 38(10): p. 9. [0136] 13. Ormiston,
J. A. and P. W. Serruys, Bioabsorbable coronary stents. Circ
Cardiovasc Interv, 2009. 2(3): p. 255-60. [0137] 14. Garg, S., C.
Bourantas, and P. W. Serruys, New concepts in the design of
drug-eluting coronary stents. Nat Rev Cardiol, 2013. 10(5): p. 12.
[0138] 15. Yang, J., et al., Haemo-and cytocompatibility of
bioresorbable homo- and copolymers prepared from 1,3-trimethylene
carbonate, lactides, and epsilon-caprolactone. J Biomed Mater Res
A, 2010. 94(2): p. 396-407. [0139] 16. Yang, J., et al., Hydrolytic
and enzymatic degradation of poly(trimethylene
carbonate-co-d,l-lactide) random copolymers with shape memory
behavior. European Polymer Journal, 2010. 46(4): p. 783-791. [0140]
17. Subbu S Venkatraman, Lay Poh Tan, Joe Ferry D Joso, Yin Chiang
Freddy Boey, Xintong Wang., Biodegradable stents with elastic
memory. Biomaterials, 2006. 27(32): p. 5. [0141] 18. Umeda., K. I.
a. N., Rapid prototyping in Biomedical Engineering, Advanced
Applications of Rapid Prototyping Technology in Modern Engineering.
2011. [0142] 19. Rengier, F., A. Mehndiratta, H. Tengg-Kobligk, C.
M. Zechmann, R. Unterhinninghofen, H.- and a. F. L. G. U. Kauczor,
3D printing Based on Imaging Data: Review of Medical Applications.
International Journal of Computer Assisted Radiology and Surgery,
2010. 5(4): p. 6. [0143] 20. Melgoza, E. L., Guillem Vallicrosa,
Lidia Sereno, Joaquim Ciurana, and Ciro A. and Rodriguez, Rapid
Tooling Using 3D Printing System for Manufacturing of Customized
Tracheal Stent. Rapid Prototyping Journal, 2013. 20(1): p. 10.
[0144] 21. Lim, C. S., P. Eng, S. C. Lin, C. K. Chua, and Y. T.
Lee., Rapid Prototyping and Tooling of Custom-made Tracheobronchial
Stents. The International Journal of Advanced Manufacturing
Technology, 2002. 20(1): p. 5. [0145] 22. Park, S. A., Sang J. Lee,
Kyung S. Lim, In H. Bae, Jun H. Lee, Wan D. Kim, Myung H. and a.
J.-K. P. Jeong, In vivo evaluation and characterization of a
bioabsorbable drug-coated stent fabricated using a 3D printing
system. Materials Letters, 2015. 141: p. 4. [0146] 23. Kun Sun, K.
S., Qimao Feng, Slide fastener bioabsorbable stent and application
thereof. US20130226277 A1, 2013. [0147] 24. Sun, C., et al.,
Projection micro-stereolithography using digital micro-mirror
dynamic mask. Sensors and Actuators A: Physical, 2005. 121(1): p.
113-120. [0148] 25. Baker, E., et al., Microstereolithography of
Three-Dimensional Polymeric Springs for Vibration Energy
Harvesting. Smart Materials Research, 2012. 2012: p. 1-9. [0149]
26. Yang, J., et al., Synthesis and evaluation of poly(diol
citrate) biodegradable elastomers. Biomaterials, 2006. 27(9): p.
1889-98. [0150] 27. Serrano, M. C., L. Carbajal, and G. A. Ameer,
Novel biodegradable shape-memory elastomers with drug-releasing
capabilities. Adv Mater, 2011. 23(19): p. 2211-5. [0151] 28. Yang,
J., et al., A thermoresponsive biodegradable polymer with intrinsic
antioxidant properties. Biomacromolecules, 2014. 15(11): p.
3942-52.
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