drug-eluting bioresorbable stents for various …meitalz/articles/d2.pdf · 2016. 5. 3. ·...

ANRV281-BE08-05 ARI 12 June 2006 17:11

Drug-Eluting Bioresorbable Stentsfor Various ApplicationsMeital Zilberman1 and Robert C. Eberhart2,3

1Department of Biomedical Engineering, Faculty of Engineering, Tel Aviv University, Tel Aviv 69978,Israel; email: [email protected] Engineering Program, University of Texas Southwestern Medical Center at Dallas andUniversity of Texas at Arlington, Dallas, Texas 753903Department of Surgery, University of Texas Southwestern Medical Center at Dallas, Texas 75390

Annu. Rev. Biomed. Eng.2006. 8:153–80

The Annual Review ofBiomedical Engineering isonline atbioeng.annualreviews.org

doi: 10.1146/annurev.bioeng.8.013106.151418

Copyright c© 2006 byAnnual Reviews. All rightsreserved

1523-9829/06/0815-0153$20.00

Key Words

stent, bioresorbable polymer, drug delivery, protein delivery

AbstractA stent is a medical device designed to serve as a temporary or permanent internalscaffold to maintain or increase the lumen of a body conduit. Metallic coronary stentswere first introduced to prevent arterial dissections and to eliminate vessel recoil andintimal hyperplasia associated with percutaneous transluminal coronary angioplasty.The stent application range has expanded as more experience was gained, and en-couraging results have been obtained in the treatment of vascular diseases. Stentsare currently used for support of additional body conduits, including the urethra,trachea, and esophagus. The rationale for bioresorbable stents is the support of abody conduit only during its healing process. The stent mass and strength decreasewith time, and the mechanical load is gradually transferred to the surrounding tissue.Bioresorbable stents also enable longer term delivery of drugs to the conduit wallfrom an internal reservoir and abolish the need for a second surgery to remove thedevice. The present review describes recent advances in bioresorbable stents, focus-ing on drug-eluting bioresorbable stents for various applications. Controlled releaseof an active agent from a stent can be used to enhance healing of the surroundingtissues, to increase the implant’s biocompatibility, as well as to help cure certain dis-eases. Because a lot of research in this field has been done by us, examples for thesefunctions are described based mainly on developments in our laboratories.

153

Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

INTRODUCTION: MEDICAL SUPPORT DEVICES(STENTS) FOR VARIOUS APPLICATIONS

Coronary Stents

Coronary stenting has become an established mode of treatment in percutaneoustransluminal coronary interventions. It has been shown to reduce late resteno-sis relative to conventional balloon angioplasty (1–3). Early designs, including theWallstent (Schneider), Palmaz-Schatz ( Johnson & Johnson), Wiktor (Medtronic),and Gianturco-Roubin (Cook) stents have been replaced by the Micro (AVE),Multilink (ACS), and others (4–7). The metals used to prepare these stents are selectedfor strength, elasticity, and malleability or shape memory. Stainless steel, tantalum,and nitinol alloys are among the most commonly used materials (6–9). Nitinol offerssuperelastic and thermal shape memory properties, which allow self-expansion ofthe stent during deployment and thermally induced collapse for theoretical removalprocedures (10).

The incidence of restenosis remains high despite technical and mechanical im-provements. This restenosis is a result of in-stent neointimal hyperplasia causedby proliferation and migration of vascular smooth muscle cells (VSMCs) inducedby vessel wall injury (11). The pathology of restenosis stems from a complexinteraction between cellular and acellular elements of the vessel wall and theblood (12). Some antiproliferative and antiinflammatory agents have been shownto elute slowly from polymer coatings and to be associated with reduced neoin-timal formation in animal models. Two antiproliferative agents, paclitaxel (13,14) and sirolimus (15), have been used in humans with promising preliminaryresults.

Paclitaxel is a natural or semisynthetic diterpene composed of a rigid texanering and a flexible side chain. It is an antineoplastic agent (16, 17). However, itis potentially cardiotoxic, and the dose of paclitaxel that can be delivered safetyhas yet to be resolved (18, 19). Paclitaxel has been shown to markedly attenu-ate stent-induced intimal thickening of the lumen (20, 21). Paclitaxel’s antiprolif-erative effect is reversible (22). Its short cellular residence time (1 h), along withthe reversible antiproliferative activity, suggests that it should be formulated insustained-release dosage form (23). Sirolimus (rapamycin) is a carbonyl lactone-lactam macrolide that has been shown to inhibit VSMC growth. This inhibitionhas been reported to be concentration dependent, with a threshold limit of 16.7ng/ml (24). Sirolimus has been shown to be effective with a remarkable restenosisrate of almost 0% (25–27). However, some criticism has been expressed regard-ing the absence of data in complex lesions, as well as long-term data (28). Resultsof the RAVEL and SIRIUS trials demonstrated that sirolimus-eluting stents effec-tively inhibit restenosis in humans (29, 30). The TAXUS trials revealed significantinhibition of coronary stenosis by paclitaxel (31–34). Drug-eluting polymer-coatedstents have thus moved into the limelight as vehicles for local drug administration(35).

154 Zilberman · Eberhart

Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Other Uses for Stents

The range of stent applications has expanded as more experience has been gainedand following encouraging results in the treatment of vascular diseases. Stents havebeen used for treating urethral obstructions caused by benign prostatic hyperplasiaand for treating benign or malignant tracheobronchial obstructions. They have alsobeen used for supporting the neonatal trachea in tracheal malacia; for treating benignand malignant esophageal, gastrointestinal, and bile duct strictures; and for treating(stents and stent-grafts) arterial dissections, aneurysms, and various neurovasculardiseases.

Stents in urology. Stents have been used to prevent urine retention following ther-mal treatment of benign prostatic hyperplasia (BPH) by various means, includingtransurethral microwave therapy and direct vision laser ablation of the prostate.Several stent designs were shown to prevent obstruction of the prostatic urethraand restricture of the anterior urethra. These stents include the Barnes, Finnishbiodegradable self-reinforced polyglycolic acid (SR-PGA) spiral; the Nissenkorn;and the Trestle stents (36–38). In clinical studies, researchers have used biodegrad-able stents to treat benign prostatic hyperplasia. The results obtained yield morepositive outcomes compared with those using suprapublic catheters (39–43). Self-reinforced poly(l-lactic acid) (PLLA) bioresorbable spiral stents are also undergoingevaluation for use in the anterior and posterior urethra and in the upper urinary tractfor preventing urinary retention and for repairing local urethral trauma or defects(44, 45).

Stents for management of tracheobronchial obstruction. Tracheobronchial ob-struction owing to either benign or malignant disease causes significant morbidityand mortality. Metal stents, which were originally developed for the vascular system,have been adapted for lesions involving the tracheobronchial tree and include theGianturco-Z (William Cook Europe), Palmaz (Johnson & Johnson), Strecker (BostonScientific), Ultraflex (Boston Scientific), and Wallstent (Boston Scientific) stents (46).These stents were used successfully for treating patients with bronchogenic cancer,inoperable esophageal tumors, primary tracheal tumors, and metastatic malignancies.Bioresorbable external tracheal stents have been investigated for treating pediatrictracheal malacia, for solving the problem of limited tracheal growth in children withrigid external fixation, and for avoiding the necessity of a second procedure for re-moving the synthetic material (46–48). Metal stents and nonexpandable tubular stable(nondegradable) polymeric stents were tried as internal stents to treat infants withtracheomalacia (49, 50). The results from these studies suggest that stenting is apromising method for treating tracheal obstruction.

Stents in the esophagus and gastrointestinal tract. Many malignant and benignesophageal and gastrointestinal strictures can be treated by minimally invasive alter-natives to surgery, including the use of stents. The most commonly used stents in

www.annualreviews.org • Drug-Eluting Bioresorbable Stents 155

Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

the esophagus and in gastrointestinal tract are the Esophacoil (Instent), Flamingo(Boston Scientific), Gianturco-Z (William Cook Europe), Ultraflex (Boston Scien-tific), and Wallstent (Boston Scientific) stents. These stents have generally been shownto be effective in relieving esophageal dysphagia (51–53), a success that has led to theemployment of stents to manage lesions of the gastrointestinal tract, including thestomach, pylorus, duodenum, upper small intestine, and colon (51, 52). Use of biore-sorbable materials for the esophageal stent is currently being explored.

BIORESORBABLE POLYMER STENTS

Bioresorbable polymeric stents have attracted much attention as alternatives to metal-lic stents. There are several reasons for fabricating a stent composed of a biodegrad-able polymeric material. Bioresorbable polymeric vascular stents have the potentialto remain in situ for a predicted period of time, keeping the vessel wall patent andthen degrading to nontoxic substances. Accumulating evidence indicates that the useof a bioresorbable coronary stent dramatically decreases the need for a prosthesisafter six months (54). Bioresorbable stents are preferable for treatment of tracheoma-lacia in newborns and infants because removal surgery is not necessary. Furthermore,bioresorbable stents can be used as support devices as well as platforms for drug andprotein delivery to the conduit wall in all of the above-mentioned applications.

Bioresorbable Polymers Used in Polymeric Stents

PLLA, PGA, poly ε-caprolactone (PCL), and poly-d,l-lactic acid (PDLLA) are themost frequently used aliphatic poly(α-hydroxy-acids) for preparing bioresorbablestents (55–57). The semicrystalline PLLA and PGA have high initial tensile strength,permitting a robust mechanical design. The PLLA total degradation time is approx-imately 24 months, whereas that of PGA is 6–12 months (Table 1). PLLA is oneof the most important biodegradable polymers, and is used in a wide range of clini-cal applications, including devices for orthopedic (58–60) and cardiovascular surgery(61), sutures (62), and drug-delivering implants (63, 64). This polymer is a very goodchoice, especially for the first three above-mentioned applications, where high me-chanical strength and toughness are required. PLLA can be formed into fibers, films,tubes, and matrices using standard processing techniques such as molding, extrusion,spinning, and solvent casting (62). PCL is also a semicrystalline polymer with a rel-atively high degree of crystallinity (similar to PLLA and PGA). However, it exhibitslower strength and modulus than PLLA and PGA owing to its low glass transitiontemperature (below room temperature, see Table 1). PDLLA is actually a randomcopolymer that consists of l-lactic acid and d-lactic acid monomers. It is thereforeamorphous and cannot exhibit crystalline structures. Its strength and modulus arelower than those of PGA and PLLA. Polydioxanone (PDS) has gained increasinginterest in the medical and pharmaceutical fields owing to its excellent biocompat-ibility (62). Although it is a semicrystalline polymer, it also exhibits lower strengththan PLLA and PGA because it has a low glass transition temperature, similar toPCL. We are the first research group to investigate stents made of the PDS polymer.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Table 1 Characteristics of typical bioresorbable polymersa

PolymerMeltingpoint (◦C)

Glass transitiontemperature (◦C)

Modulus(Gpa)

Degradationtime (months)

PGA 225–230 35–40 7.0 6–12PLLA 173–178 60–65 2.7 >24PDLLA Amorphous 55–60 1.9 12–16PCL 58–63 (−65)–(−60) 0.4 >24PDS N/A (−10)–0 1.5 6–1285/15 PDLGA Amorphous 50–55 2.0 5–675/25 PDLGA Amorphous 50–55 2.0 4–550/50 PDLGA Amorphous 45–50 2.0 1–2

aAdapted from “Synthetic Biodegradable Polymers as Medical Devices,” MPB Archive, March1998.PGA: poly(glycolic acid); PLLA: poly(l-lactic acid); PDLLA: poly(DL-lactic acid); PCL:poly(ε-caprolactone); PDS: polydioxanone; PDLGA: poly(DL-lactic-coglycolic acid).

The main results found for these stents are presented in Drug and Protein-LoadedBioresorbable Polymer Stents, below.

Useful combinations of the above-mentioned materials (copolymers and blends)can be created to alter the mechanical properties and drug-release profiles of bioac-tive agents from polymeric structures based on these polymers. For example, 10/90PDLGA, a random copolymer that contains 90% glycolic acid and 10% lactic acid,has relatively high strength but is more flexible than PGA and degrades faster. OtherPDLGA formulations that contain relatively high lactic acid contents, such as 85/15PDLGA, 75/25 PDLGA, and 50/50 PDLGA, are amorphous. Therefore, they donot exhibit high strength and modulus. However, they can be used for support incases that do not require high strength, such as in neural stents.

All above-mentioned materials degrade principally by simple hydrolysis of theester bond in the polymer backbone. Partial chain scission degrades the polymer to10–40 μm particles. These particles can be phagocytized and metabolized to carbondioxide and water, which are, of course, fully resorbed. The polymer’s degradationtime is a function of its chemical structure and molecular weight. Crystallinity con-tributes to a higher degradation time because crystalline domains are denser thanamorphous domains and water molecules cannot penetrate them easily. Table 1demonstrates that a relatively small number of polymers provides a large varietyof degradation times for various medical support applications.

Bioresorbable Stents for Various Applications

Several bioresorbable stents for various applications have been reported (65, 66).Kemppainen et al. (65) and Multanen et al. (67) reported spiral PLLA stents, andBrauers et al. (66) described tubular PDLLA and poly(DL-lactic–co–glycolic acid)stents, both for urethral applications. None of these stents is expandable.

Several early designs of expandable bioresorbable stents have been developed asalternatives to metallic vascular stents (68–75). The first biodegradable stent was


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

developed by Stack & Clark of Duke University in the early 1980s (68, 69). They in-vestigated several biodegradable polymers and chose PLLA as the stent material. ThisDuke University stent was constructed of a specialized PLLA polymer woven into adiamond-braided pattern from eight polymeric strands. It was designed to withstandcompression pressures of up to 1000 mm Hg (the Palmaz-Schatz stent can with-stand pressures of 300–500 mm Hg and maintain its radial strength for one month).In vivo studies demonstrated minimal thrombosis and inflammatory responses, andmoderate neointimal growth. Gao et al. (70) reported the Tianjin/Beijing stent, aPDLLA/PCL stent with an inner heparin layer, deployed with a balloon catheter andemployed heating and pressurization. This stent produced mild neointimal prolifer-ation in swine carotid artery models at two months. Yamawaki et al. (71) reported aKyoto University PGA coil stent, which exhibited thrombus deposition in canine im-plant studies, but no subacute closure. Nuutinen et al. (74) developed knitted PLLAand 80/20 PDLGA stents with mechanical properties similar to those of commer-cially available metallic stents. They found that the knitting geometry has a markedeffect on the stents’ mechanical properties. Such knitted bioresorbable stents are moresuitable for urological, gastrointestinal, and tracheo-bronchial indications, where thesize of the delivery device is not critical, as opposed to peripheral vascular or evencoronary indications, for which a small-diameter delivery device is critical. Saito et al.(75), for example, studied knitted stents in a rabbit trachea model.

DRUG AND PROTEIN LOADED BIORESORBABLEPOLYMER STENTS

Fiber-Based Stents

Drug-eluting fiber-based stents have been developed mainly for blood vessel support,but the stent designs can be used also for ureteral support.

Vascular stents. Tamai et al. (72, 76) described the Igaki/Tamai stent, a biore-sorbable balloon expandable zigzag coil design based on a PLLA monofilament.The thickness of the stent strut is 0.007 inches (0.17 mm) and the stent surface areais 24% at an arterial diameter of 3.0 mm. The stent has a self-expanding capacity,with an expansion range of up to 4.5 mm. This bioresorbable stent combines the fea-tures of a thermal self-expandable and a balloon expandable stent. The stent initiallyautoexpands in response to the heat transmitted by a delivery balloon inflated witha 70◦C contrast-water mixture (50◦C at the balloon site). Subsequent expansion isobtained by inflation at a moderate to high pressure (6 to 14 atm). This stent willcontinue to expand to its nominal size within the following 20 to 30 min at 37◦C. Themost important and unique innovation made by Tamai et al., compared with priorinvestigators, is the change in stent design from a knitted pattern to a zigzag helicalcoil design.

The success of bioresorbable polymeric stents depends not only on the biocom-patibility of the stent material but also on the ability of the manufactured stent itself,including the stent design, the scaffolding strength, the biodegradation period, etc.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Vessel wall injury causes neointimal proliferation (77), and the severity of vesselinjury is strongly correlated with neointimal thickness and the percentage of stenosisafter balloon angioplasty or stenting (78). The stent design may reduce the extentof vessel wall injury caused by the stent implantation and may influence the neointi-mal proliferation and the inflammatory response. Animal studies of the Igaki/Tamaistents have demonstrated that the PLLA coil stent reduced the percentage of steno-sis in porcine coronary arteries at 2 weeks from 64% to 19%, compared with thePLLA knitted type stent (79). Minimal neointimal hyperplasia was found within thePLLA coil stents, whereas moderate to severe neointimal hyperplasia was observedin knitted PLLA stents. The PLLA coil stents also exhibited long-term biocompat-ibility with a minimal inflammatory response in porcine coronary arteries after 16weeks (80, 81). Unlike stents used in previous studies, the Igaki/Tamai stent was madefrom a high-molecular-weight PLLA that resulted experimentally in a minimal in-flammatory response, compared with that observed in previous reports. In this study,25 stents were deployed successfully in 15 patients. No stent thrombosis or majorcardiovascular events occurred within 30 days. After 6 months, restenosis occurredin 10.5% of the patients and target lesion revascularization occurred in 6.7%. Thisstudy actually provided the first report on immediate and 6-month results followingimplantation of a bioresorbable PLLA stent in humans.

Tranilast [(N-(3,4-dimethoxycinnamoyl) anthranilic acid] is an antiallergic drugthat also inhibits the migration and proliferation of VSMCs induced by platelet-derived growth factor and transforming growth factor β1 (12, 82). Tranilast wasfound to significantly suppress vascular intimal hyperplasia in rabbit artery afterballoon injury (83, 84) and in pig coronary arteries after balloon angioplasty andstent implantation (85). The tranilast-eluting Igaki-Tamai stent was made of a PLLAmonofilament mixed with tranilast, and its shape was similar to that of the regularIgaki-Tamai stent without the drug. The radial compression strength of the drug-loaded stent was approximately 10% lower than that of the neat stent.

Drug-mixed polymer stents can be loaded with larger amounts of drug than candrug-coated stents because the polymer stent struts can contain the drug. Tsuji et al.(35) compared the tranilast content in the tranilast-eluting Igaki-Tamai stent with thatin a tranilast-coated Palmaz-Schatz stent that was coated with a 20–40 mm layer oftranilast-mixed PCL. The tranilast content of the former was found to be at least fourtimes higher than that of the latter. The authors developed three types of tranilast-eluting stents: a noncoated stent that can release tranilast rapidly, a PCL-coated stentfor relatively slow release of tranilast through the PCL barrier layer, and a stent witha tranilast-loaded PCL coating designed to rapidly release the drug followed by aslower release through the PCL layer. The authors stated that clinical experimentsare necessary to elucidate these stents’ safety and efficacy.

Uurto et al. (86) evaluated in vivo a new fiber-based drug-eluting bioresorbablevascular stent, with respect to biocompatibility, neointimal hyperplasia formation,and reliability. Monofilaments made of a polymer consisting 96% l-lactic acid and4% d-lactic acid were manufactured by melt spinning. The stents were formed froma mesh design consisting of 16 monofilaments braided over a mandrel. These stentswere coated 50/50 with two bioactive agents: dexamethasone and simvastatin. Both


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

PCL and a polymer containing 50% l-lactic acid and 50% d-lactic acid were used ascoating materials. The in vivo results (pig experiments) indicated that all stented ar-teries were angiographically patent. The dexamethasone-loaded stents decreased themean luminal diameter compared to other stents and they induced minimal intimalhyperplasia. The vascular injury scores demonstrated only mild vascular trauma forall studied stents. Hence, the authors concluded that bioresorbable polymer stentsappear to be biocompatible and reliable, and can be used as a local drug deliveryvehicle. The findings showed a need for further investigation to prove the efficacyand safety of this new biodegradable drug-eluting stent.

The two fiber-based drug-eluting bioresorbable vascular stents described abovedemonstrated promising results. Other fiber-based vascular stents were developed bycompanies, and their studies are in progress. The growing number of such stents in-clude, for example, the paclitaxel-loaded REVA stent (developed by REVA Company)and a stent developed by TissueGen and Endovasc companies.

PLLA multiple-lobe vascular stents. A fiber-based expandable stent design hasbeen developed and studied in our laboratory (56, 87, 88). This stent is preparedusing a linear, continuous coil array principle, by which four furled lobes convertto a single large lobe upon balloon expansion (Figure 1). Melt-extruded fibers witha 0.15 mm diameter were woven continuously around a four-mandrel array (onecentral, three peripheral) into a four-lobe configuration. Three longitudinal fiberswere interwoven and glued to the coil for mechanical support. After fabrication, a

Figure 1The design concept of the vascular fiber-based stent: (a) predilated; (b) predilated, side view;(c) dilated; and (d) dilated, side view.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

conventional angioplasty balloon catheter is inserted in the central lobe and the stentcan be deployed at the target site. The fully expanded stent has a helical coil structurewith three longitudinal reinforcing fibers (Figure 1). The stents’ initial and finaldiameters are adjustable.

The bioresorbable fibers were made from a relatively high-molecular-weightPLLA (RESOMERTM L21, Boehringer Ingelheim, Germany) with an inherent vis-cosity of 3.6 dL/g in CHCl3 at 30◦C. This polymer was melt spun at 190◦C and drawnat 80◦C to draw ratios of 3:1 to 8:1, to create fibers with good tensile mechanical prop-erties. The undrawn fiber is weak and almost cannot be handled. The yield stress,ultimate tensile strength, and young modulus increased dramatically as the draw ratioincreased and the maximal strain decreased, i.e., the fiber became stronger, tougher,and more brittle. These changes occur mainly due to the structural changes that leadto an increase in the polymer’s degree of crystallinity (87). The 8:1 drawn fibers ex-hibited the highest tensile strength (980 MPa) and modulus (4.9 GPa) together withgood ductility (50% strain). Most of the studied stents were therefore prepared fromthese fibers.

Stents with a length of 15 mm, a final (dilated) diameter of 3.0 mm, and a predilateddiameter of 1.8 mm were used in this study. Most of the study focused on single-fiberstents. However, PLLA double-fiber stents were also investigated. Each dilated coilstent contained 12 loops, each bonded to three longitudinal support fibers, i.e., 36binding points per single-fiber stent and 72 binding points per double-fiber stent.The initial radial compression strength of the dilated form of both types of stentsexceeded 200 KPa. It should be mentioned that the maximal pressure that can beapplied using our radial compression chamber is 200 KPa (87). Higher pressurescould therefore not be measured.

The stents were immersed in PBS at 37◦C and samples were removed periodi-cally to investigate the effect of in vitro degradation on their mechanical properties.The mode of failure observed was rupture of binding points, where the longitudi-nal support fibers were glued to the coil. The radial compression pressure neededto create a rupture of at least one binding point as a function of degradation timeis presented in Figure 2. The number of ruptures (binding points that failed) foreach degradation time is indicated in parentheses. The PLLA stents were studiedfor 20 weeks. Both single- and double-fiber types of stents did not undergo any fail-ure throughout the first eight weeks of exposure to an aqueous medium. Afterward,the radial compression pressure required to create a rupture at the binding pointsexhibited a linear decrease with time and the number of ruptures increased withtime (Figure 2). Because the double-fiber design has more binding points than thesingle-fiber design, each binding point of the former is exposed to a smaller pres-sure. Thus, a higher total pressure is required to fail the double-fiber stent. Thesestents generally exhibited good radial compression endurance. They resisted at least150 KPa (approximately 75% of the initial strength) and exhibited only few rupturepoints, whereas most of the binding points remained intact. The combination of thesuggested design and the relatively high-molecular-weight PLLA may thus be appli-cable for supporting blood vessels for at least 20 weeks and was chosen for furtherstudies.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

160

180

200

220

80 4 12 16

Single fiber stent

Double fiber stent

20Ra

dia

l c

om

pre

ss

ion

str

en

gth

(k

Pa

)

Time (weeks)

100

120

140

(0) (0) (0) (0)

(2)

(3)(1)

(3)

(4)

Figure 2The PLLA stent redial compression strength as a function of immersion time in PBS at 37◦C:single- (� ), and double-fiber (•) stents. The number of rupture points is indicated under eachdata point in parentheses.

Bioresorbable stents can simultaneously support blood vessels and serve asdrug/protein delivery platforms. Certain drugs, such as steroids, can be incorporatedinto the PLLA fiber during melt processing. However, only small drug quantities(less than 5% w/w) can be incorporated in dense polymeric structures such as fiberswithout having an adverse effect on the mechanical properties of the fiber and thestent. Furthermore, most drugs and all proteins are destroyed when exposed to thehigh melt processing temperature. To solve this problem, we developed and studiednovel drug-loaded composite fibers (87, 89–92) and stents prepared from these fibers.These unique fibers actually resemble sutures (Figure 3a) and contain two sections:(a) a dense core fiber with good tensile mechanical properties (high strength and goodflexibility), such as the fibers that were described here, and (b) a porous shell, i.e., abioresorbable coating that contains the drug molecules. This section is prepared viaemulsions using mild processing conditions. Two types of such core/shell fiber struc-tures were developed and studied: a shell made of a highly porous foam-like structure[prepared using the single emulsion technique (89, 90, 92)] as presented schematicallyin Figure 3b, and a shell made of microspheres [prepared using the double emulsiontechnique (87, 91)] as presented schematically in Figure 3c.

Protein-loaded bioresorbable microspheres were developed and bound to thePLLA fibers and stents. These microsphere reservoirs, which were prepared bythe double emulsion technique, can be loaded with biologically active aqueous ornonaqueous molecules. Because mild materials and processing steps are used, thesemicrospheres can be loaded with all drugs, proteins, and gene transfer vectors. Theprocessing conditions can be controlled to yield single-reservoir or multiple-reservoirmicrospheres, as shown in Figure 4. The microspheres are well attached to the


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Figure 3Composite drug-loaded core/shell fiber structures: (a) general view (photograph) of a fiber,(b) a schematic representation showing a dense core fiber coated with drug/protein loadedporous shell, and (c) a schematic representation showing a dense core fiber coated with amicrosphere shell.

fibers, as seen in Figure 4a. The initial radial compression strength of all types ofmicrosphere-loaded stents exceeded 200 KPa. This indicates that the partial disso-lution of the surface layers that was performed to enable microsphere attachmentto the fiber had practically no affect on the strength of the stent (87). It has beendemonstrated that the microsphere structure strongly affects the release profile ofthe agent from the microspheres and from the microsphere-loaded stents (87, 91).


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Figure 4SEM micrographs of (a) microsphere-loaded PLLA fiber, (b) single-reservoir PDLGAmicrosphere with thin wall, (c) single-reservoir PDLGA microsphere with thick wall, (d)multiple-reservoir PDLGA microsphere, and (e) matrix-like PDLGA microsphere withnumerous small reservoirs inside.

In one of our experiments, four types of microspheres were prepared frompoly(DL-lactic-coglycolic acid) consisting of 75% DL-lactic acid and 25% glycolicacid (75/25 PDLGA). Two of these microsphere types had a relatively low initialmolecular weight (inherent viscosity 0.35 dL/g, ∼43,000 Daltons) and two had amedium initial molecular weight (inherent viscosity 0.69 dL/g, ∼118,000 Daltons).Each of these two types of microspheres was prepared twice, once with a thin wall andonce with a thick wall, as presented in Figures 4b and 4c, respectively. The micro-spheres’ wall thickness can be controlled by choosing the proper relative quantities ofwater and polymer phases of the double emulsion (87). Albumin, an inexpensive pro-tein, was encapsulated in the microspheres. Albumin was chosen to serve as a modelfor controlled release of water-soluble agents such as certain drugs and proteins as wellas viral gene transfer vectors from the stent. Most of the albumin was found to be en-capsulated in these 10–70 μm microspheres (mean efficiency of 85%). The cumulative


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Several reservoirs

6

Pro

tein

cu

mu

lati

ve

re

lea

se

(%

)

Time (weeks)

a

0

20

40

60

80

100

0 4 531 2 60 80

Time (days)

0 4020

b

Low MW thin wall

Low MW thick wall

Regular MW thin wall

Regular MW thick wall

Many small reservoirs

Figure 5In vitro cumulative protein release from microsphere-based structures. (a) Effect ofmicrospheres’ initial molecular weight and wall thickness on the release profile from stentsloaded with single-reservoir microspheres. (b) Effect of microsphere structure on releaseprofile from fibers loaded with multiple-reservoir microspheres.

albumin release profiles from the four types of microsphere-loaded stents are pre-sented in Figure 5a. In general, a burst effect is accompanied by linear cumulativerelease profile, as expected for a single-reservoir system, such as these microspheres.The microspheric single reservoirs enable effective protein release from the stents.The drug release rate increased with the decrease in the microsphere’s initial molec-ular weight owing to a higher degradation rate. Furthermore, the thick microsphereshell was found to be effective in reducing the burst effect. Multireservoir micro-sphere types (Figure 4d,e) are currently being studied, in addition to drug releasefrom stents loaded with single-reservoir microspheres. It has been demonstratedthat a low mixing rate of the internal emulsion results in relatively small micro-spheres with numerous small reservoirs inside (91), as presented in Figure 4e. Thisunique matrix-like structure, which was obtained owing to a reduced solvent evapo-ration rate, enabled a relatively low burst release and a more moderate release profile(Figure 5b).

Current research focuses mainly on the development of a method for incorpo-rating non-water-soluble drugs, such as paclitaxel, in this stent design. As explainedin the introduction (see Coronary Stents, above), paclitaxel is an antiproliferativedrug that can prevent restenosis by inhibiting vessel smooth muscle cell proliferation(93). Effective controlled release of the drug molecules from the stent to the bloodvessel wall is desired. Slow-release paclitaxel applied perivascularly totally inhibitsintimal hyperplasia and prevents lumenal narrowing following balloon angioplasty.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Studies indicate the need for controlled paclitaxel release owing to its narrow toxic-therapeutic window and highly hydrophobic nature (94). The results presented hereindicate that the microspheric reservoirs enable effective drug release from the stents.However, because the release rate of the described systems was relatively slow andthe chosen active agent (paclitaxel) is not water soluble, we decided to focus on thedevelopment and study of highly porous bioresorbable fiber structures, as describedin Figure 3b, that will be used to build stents. The shell section of such fibers isexpected to degrade much faster than the microspheres, and as a result they willrelease a non-water-soluble drug within 2–4 weeks, as desired for this application(95, 96). SEM micrographs showing a general view and a cross section of such fibersare presented in Figure 6a,b. This research is in progress, and it has already been

Figure 6SEM micrographs of a core/shell fiber structure consisting of a PLLA core fiber coated with aporous PDLGA structure: (a) general view, (b) part of a cross section.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

found that the paclitaxel release profile from such fibers generally exhibits a releaseprofile typical for diffusion-controlled systems (92). The active agent’s content andthe polymer content of the emulsion’s organic phase demonstrated a significant effecton both the shell microstructure and on the release profile, whereas the emulsion’sorganic:aqueous phase ratio affected these fiber characteristics only in certain cases.

Bioresorbable multilobe stents based on other polymers for short-term ap-plications. Procedures to remove kidney stones can be extremely painful and ofteninclude the insertion of a nondisposable plastic or ruber ureteral stent to prevent fur-ther damage of the urinary tract and promote healing. Presently, most ureteral stentsrequire a second (cystoscopy) procedure for removal. By inserting a bioresorbablestent at the time of the first procedure, one could eliminate the pain and expense ofthe second procedure. At the present time, there are no bioresorbable urinary stentsavailable in the market. Such stents should degrade much faster than PLLA ones.

Our PLLA stents demonstrated excellent initial strength and strength retentionfor 20 weeks. Because the PLLA’s degradation rate is relatively slow, other polymersshould be considered for moderate and short-term applications, such as in urinaryapplications. We therefore also investigated polydioxanone (PDS) and poly(glycolic-coε-caprolactone) (PGACL), which exhibited moderate and fast degradation rates,respectively (97).

PDS has gained increasing interest in the medical and pharmaceutical fields ow-ing to its excellent biocompatibility (62). It has been introduced into the market asa suture and as a bone pin [ORTHOSRB (98, 99)]. Its total degradation time is 6–12 months. PGACL is a block copolymer of glycolide and ε-caprolactone. This uniquebioresorbable polymer degrades completely within 3–4 months. It offers reducedstiffness compared with pure poly(glycolic acid) and has been introduced into themarket by Ethicon as a monofilament suture (Monocryl). The initial tensile mechan-ical properties of these two fibers compared with those of PLLA, as measured byus, are presented in Table 2. PLLA demonstrated a very high tensile strength andmodulus and a moderate ultimate strain (ductility), whereas PDS exhibited moderatetensile strength and modulus and a relatively high ductility, and PGACL exhibitedhigh tensile strength and ductility and moderate modulus. These three polymers arethus totally different from each other and exhibit a relatively large variety of initialmechanical properties. It should be emphasized, again, that in order to be able toproduce stents from these fibers they should combine relatively high strength withsufficient ductility and flexibility. All three fiber types exhibited these required prop-erties. The weight retention profiles of the three fibers are presented in Figure 7a.

Table 2 Tensile mechanical properties of bioresorbable fibers

Fiber Tensile strength (MPa) Modulus (MPa) Strain (%)PLLA 967 5000 50PDS 583 367 161PGACL 961 636 151


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

60

70

80

90

100

Re

ten

tio

n o

f w

eig

ht

(%)

Ra

dia

l c

om

pre

ss

ion

str

en

gth

(k

Pa

)

Time (weeks)

0

50

100

150

200

250

0 1 4 5 6

Time (weeks)

(2)

(5)

(12)

(3)

(3)

(3)

(5)

(6)(10)

b

a

(0) (0) (0) (0) (0)(0) (0)

0 1 2 3 4 5 6

2 3

PLLA

PLLA

PDS

PDS

PGACL

PGACL

Figure 7(a) Retention of weight, and (b) radial compression strength of single-fiber bioresorbable stents(number of ruptured points is indicated) as a function of degradation time in PBS at 37◦C.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

PLLA did not exhibit any weight loss following exposure to an aqueous medium dur-ing the six weeks of study, whereas PDS lost approximately 2% of its initial weightand PGACL gradually lost 30% of its initial weight within 6 weeks of immersion inan aqueous medium.

The radial compression strength as a function of degradation time for single-fiberPDS and PGACL stents compared with PLLA stents is presented in Figure 7b. ThePDS stents began losing mechanical strength after 3 weeks of degradation, but pre-served good strength for 6 weeks. In contrast, PGACL demonstrated a faster decreasein radial compression strength (the slope of the curve) that began immediately afterexposure to the aqueous medium. This study thus indicates that PLLA stents cansupport body conduits for relatively long periods of time (20 weeks) and can there-fore be used as endovascular stents, whereas PDS stents can afford good support for5–6 weeks and PGACL stents can be applied for only 2 weeks. PDS and PGACLstents are therefore utilizable for short-term applications such as local support ofthe urinary track after surgery as well as after other interventions. In these cases,drug-eluting stents can be prepared using the methods described above for core/shellcomposite PLLA fibers. Incorporation of antiinflammatory agents in these stents maybe beneficial.

Film-Based Stents

Drug-eluting film-based stents have been developed to support blood vessels and thetrachea.

Vascular stents. We demonstrated the successful transfer and expression of anuclear-localizing ß-Gal reporter gene in cells in the arterial wall of rabbits after theimplantation of biodegradable stents made of bioresorbable PLLA/PCL films creat-ing a porous tubular structure impregnated with a recombinant adenovirus carryingthat gene (100). Although we used a nonexpandable stent design, we demonstratedthe exciting possibility of transferring genes that encode for key proteins in the centralregulatory pathways of cell proliferation inside arterial wall cells, using bioresorbablestents as vehicles.

Vogt et al. (101) reported a novel PDLLA double-helical stent designed accordingto the material’s properties, based on a finite element method. This stent was man-ufactured using the controlled expansion of saturated polymers (CESP) for moldingthe PDLLA. This method is characterized by a low process temperature that enablesthe processing of thermally sensitive polymers and the incorporation of biologicallyactive substances. This stent exhibited sufficient mechanical stability and a significantpotential to reduce restenosis after vascular intervention was observed following theincorporation of paclitaxel. Paclitaxel was shown to effectively inhibit proliferationand coronary stenosis in a pig model after vascular injury in a long-term course. Theauthors suggest that in vitro pharmacokinetics with a very slow release pattern ofpaclitaxel over a time span of more than 2 months might contribute to this favorableoutcome. The authors also reported that paclitaxel loading did not increase toxicity,as reflected by media necrosis or delayed healing of endothelial cells.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Alexis et al. (23) studied the in vitro release kinetics of paclitaxel and rapamycinfrom solution-cast PDLLA and PDLGA films and from nonexpandable helical stentsprepared from strips cut from the films. The results demonstrated that the releasemechanism for both drugs combined diffusion and degradation. The films and stentsexhibited the same release profiles, indicating homogenous degradation kinetics. Noburst effect was observed for either drug. This is especially important during pacli-taxel administration, where cardiotoxicity is sometimes a concern. The authors alsodemonstrated that the release period can vary from one month to several months,depending on the matrix polymer.

PLLA tracheal stents. The metal stents currently used for blood vessel support canalso serve for airway support (49, 50). However, an ideal bioresorbable tracheal stentwould support the neonatal trachea in tracheal malacia until the airway matures, andwould then be totally resorbed, thus negating the need for a removal operation. Wehave developed a novel expandable stent prepared from a solution-cast PLLA-basedfilm (73, 102). The stent design is presented in Figure 8. Dexamethasone (DM), asteroid antiinflammatory drug, is incorporated into this film-based stent to providecontrolled local release of the drug during the mechanical support phase. In additionto its well-known antiinflammatory activity, DM has been demonstrated to inhibitthe fibrotic response. Addition of this drug may therefore be helpful in the preventionof proliferative reactions, such as airway stenosis. This support structure was testedas a tracheal stent to support the tracheal airway and to treat tracheomalacia both invitro and in vivo.

In vitro studies of the films and tracheal stents focused on their mechanical prop-erties in light of the morphology and degradation and erosion processes. The PLLAfilms generally exhibited good initial mechanical properties. Exposure of the filmsto an aqueous medium resulted in decreased tensile strength, modulus, and ductilityowing to chain scission and morphological changes. These films underwent somedegradation and small changes in morphology during the 20 weeks of study, andretained good mechanical properties (73).

Our PLLA tubular support structures (stents) developed from these films demon-strated good mechanical properties. Their initial radial compression strength (at least200 KPa) was six times higher than the minimal strength required for supporting aninfant trachea. The stent’s radial compression strength is determined mainly by thepolymer structure, whereas drug incorporation has a minor effect on the initial stentstrength. Exposure to radial compression stress results in reversible elastic defor-mation of most stents, whereas certain stents prepared from relatively brittle poly-mer films exhibited a sudden brittle fracture. These stents retained good mechanicalstrength for at least 20 weeks in an aqueous medium at 37◦C (73). Incorporationof DM in the PLLA stents contributes to improved strength retention owing tothe drug’s stiffness and good dispersion. Preliminary in vivo studies of these stentsdemonstrated excellent proof of principle and tolerable biocompatibility.

Drug release from films used in tracheal stents. A method developed by us forcontrolling drug location/dispersion in the film was reported (102) and is described


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

Figure 8The design concept of the tracheal film-based stent: (a) photograph of the dilated form, (b)schematic representation of the predilated form, and (c) schematic representation of thedilated form.

briefly in this section. Bioresorbable polymeric films containing DM were preparedusing solution processing, accompanied by a postpreparation isothermal heat treat-ment. In this process, the kinetics of drug and polymer solidification are determinedby the solvent evaporation rate and thus also the drug dispersion/location in the film.Solubility effects in the starting solution also contribute to the postcasting diffu-sion processes, and occur concomitantly to the drying step. In general, two types of


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

polymer/DM film structures were created and studied for all studied matrix polymertypes, as presented in Figure 9, for amorphous PDLGA/DM films:

1. A polymer film with large crystalline drug particles located on its surface, aspresented in the upper micrograph of Figure 9. This structure is derived froma dilute solution of both polymer and drug and was obtained using the slowsolvent evaporation rate that enables prior drug nucleation and growth on thepolymer solution surface. This skin formation is accompanied by later polymercore formation/solidification. This structure was named the A-type.

2. A polymer film with small drug particles and crystals distributed within thebulk, as presented in the lower micrograph of Figure 9. This structure is de-rived from a concentrated solution and was obtained using the fast solventevaporation rate and resulted from drug nucleation and segregation within adense polymer solution. Solidification of drug and polymer occurred concomi-tantly. For semicrystalline-based films such as PLLA/DM, the drug is locatedin amorphous domains of a semicrystalline matrix, around the spherulites. Thisstructure was named the B-type.

Figure 9Light micrographs of polymer/drug films. (a) A-type, with large drug crystals located on thesurface of the polymer film. (b) B-type, with small drug particles and crystals located within thepolymer film.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

We have studied these structures’ morphologies and formation processes exten-sively, and a detailed model describing the structuring of these films is describedelsewhere (102). Investigation of these films focused on cumulative DM release asaffected by film morphology (drug location/dispersion in the film) and degradationprocesses. The effect of the polymer’s degree of crystallinity on the DM release profilewas also studied (103).

In general, the drug location/dispersion in the film and the polymer’s weight lossrate determine the drug’s release profile from the film. However, the DM release pro-file in certain cases of polymers with high degradation rates is determined mainly bythe polymer’s degradation, and the drug location/dispersion in the film has almost noeffect on the release profile. All release profiles from A-type films exhibited an initialburst effect of approximately 30% (Figure 10a). The hydrophobic nature of DM andthe specific interactions between DM and the host polymer do not induce a higherburst effect. The second release stage from these films occurs at an approximatelyconstant rate and is determined mainly by the polymer weight loss rate. In contrast,in most of the studied systems the matrix (monolithic) nature of the B-type film en-abled release profiles that were determined mainly by the degradation profile of thehost polymer, without any burst effect (Figure 10b). It should be emphasized thatall A-type films and some of the B-type films exhibited nearly constant DM releaserates. Such release profiles are suitable for many applications, including the currentapplication that focuses on bioresorbable tracheal stents built from these films. It wasalso found that a high degree of crystallinity contributes to slower drug release rates(compared with amorphous systems), mainly at relatively low weight losses. At high

Time (weeks)

20

0

20

40

60

80

100

0 16124 8

Cu

mu

lati

ve

d

rug

re

lea

se

(%

)

Time (weeks)

ba

200 16124 8

PLLA

PDS

A-type B-type

PLLA

85/15 PLGA

85/15 PLGA

50/50 PDLGA

85/15 PDLGA85/15 PDLGA

50/50 PDLGA

10/90 PDLGA

Figure 10In vitro cumulative drug release from polymer/drug films: (a) A-type films, (b) B-type films.The host polymer is indicated.


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

weight losses, where a porous structure is created, crystallinity has a minor effect onthe rate of drug release. In addition to degradation and drug location/dispersion inthe film, the porous structure that develops with degradation also affects the drugrelease profile from the films. For example, a fibrous pore structure probably enablesmore effective drug release than a spherical structure (103).

CONCLUSIONS

This review presents an extensive overview of published studies of new stents withcontrolled release of drugs and proteins for endovascular, tracheal, and urethral appli-cations. These fiber- and film-based devices combine a desired drug release behaviorwith good mechanical properties and can therefore be used to support various bodyconduits. Stents can therefore be used as drug/protein-delivery platforms in addi-tion to their support function, releasing active agents from the stent in a desiredcontrolled manner that can be used to enhance the healing of the surrounding tis-sues, increase the implant’s biocompatibility, or even help cure certain diseases. Theeffect of fiber/film processing parameters on the microstructure and the resultingmechanical properties and release profiles must be elucidated. New designs based oncomposite fiber structures or structured films may thus advance the therapeutic fieldof supporting and healing body conduits.

ACKNOWLEDGMENTS

The authors thank the NIH (HL-53225), the Tel-Aviv University Internal Fund, andthe Slezak Foundations for supporting this research.

LITERATURE CITED

1. Fishman DL, Leon MB, Baim DS, Schatz RA, Savage MP, et al. 1994. A ran-domized comparison of coronary-stent placement and balloon angioplasty inthe treatment of coronary artery disease. N. Engl. J. Med. 331:496–501

2. Serruys PW, de Jaegere P, Kiemeneij F, Macaya C, Rutsch W, et al. 1994. Acomparison of balloon-expandable-stent implantation with balloon angioplastyin patients with coronary artery disease. N. Engl. J. Med. 331:489–95

3. Stefanidis IK, Tolis VA, Sionis DG, Michalis LK. 2002. Development in intra-coronary stents. Hell. J. Cardiol. 43:63–67

4. Carter AJ, Scott D, Rahdert D, Bailey L, de Vries J, et al. 1999. Stent designfavorably influences the vascular response in normal porcine coronary arteries.J. Invasive Cardiol. 11:127–34

5. Dumonceau JM, Deviere J. 1999. Self-expandable metal stents. Baillieres BestPract. Res. Clin. 13:109–30

6. Ozaki Y, Violaris AG, Serruys PW. 1996. New stent technologies. Prog. Car-diovasc. Dis. 39:129–40

7. Vioralis AG, Ozaki Y, Serruys PW. 1997. Endovascular stents: a ‘break throughtechnology’, future challenges. Int. J. Cardiac Imaging 13:3–13


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

8. Nicholson T. 1999. Stents: an overview. Hosp. Med. 60:571–739. Cleveland TJ, Gaines P. 1999. Stenting in peripheral vascular disease. Hosp.

Med. 60:630–3210. Barras CDJ, Myers KA. 2000. Nitinol—its use in vascular surgery and other

applications. Eur. J. Vasc. Endovasc. Surg. 19:564–6911. Ross R. 1993. The pathogenesis of atherosclerosis: a perspective for the 1990s.

Nature 331:489–9512. Tanaka K, Honda M, Kuramochi T. 1994. Prominent inhibitory effects of trani-

last on migration and proliferation of and collagen synthesis by vascular smoothmuscle cells. Atherosclerosis 107:179–85

13. Honda Y, Grube E, de la Fuente LM, Yock PG, Stertzer SH, Fitzgerald PJ.2001. Novel drug-delivery stent: intravascular ultrasound observations from thefirst human experience with the QP2-eluting polymer stent system. Circulation104:380–83

14. Liistro F, Stankovic G, Di Mario C, Takagi T, Chieffo A, et al. 2002. First clini-cal experience with a paclitaxel derivate-eluting polymer stent system implanta-tion for in-stent restenosis: immediate and long-term clinical and angiographicoutcome. Circulation 105:1883–86

15. Degertekin M, Serruys PW, Foley DP, Tanabe K, Regar E, et al. 2002. Per-sistent inhibition of neointimal hyperplasia after sirolimus-eluting stent im-plantation: long-term (up to 2 years) cinical, angiographic, and intravascularultrasound follow-up. Circulation 106:1610–13

16. Sousa JE, Costa MA, Abizaid A, Abizaid AS, Feres F, et al. 2001. Lack ofneointimal proliferation after implantation of sirolimus-coated stents in humancoronary arteries: a quantitative coronary angioplasty and three-dimensionalintravascular ultrasound study. Circle 103:192–95

17. Sollot SJ, Cheng L, Pauly RR, Jenkings GM, Monticone RE, et al. 1995. Taxolinhibits neointimal smooth muscle cell accumulation after angioplasty in therat. J. Clin. Interv. 95:1869–76

18. Am. Soc. Health Syst. Pharm. 1989. AHFS Drug Information, pp.1075–86.Bethseda, MD: Am. Soc. Health Syst. Pharm.

19. Rowinsky EK, Donehower RC. 1995. Drug therapy: paclitaxel (Taxol). N. Engl.J. Med. 332:1002–14

20. Yamawaki T, Shimokawa H, Kozai T, Miyata K, Higo T, et al. 1998. Intralu-minal delivery of a specific tyrosine kinase inhibitor with biodegradable stentsuppresses the restenotic changes of the coronary artery in pigs in vivo. J. Am.Coll. Cardiol. 32:780–86

21. Drachman DE, Edelman ER, Seifert P, Groothuis AR, Bornstein DA, et al.2000. Neointimal thickening after stent delivery of paclitaxel: change in com-position and arrest of growth over six months. J. Am. Coll. Cardiol. 36:2325–32

22. Suh H, Jeong B, Rathi R, Kim SW. 1998. Regulation of smooth muscle cell pro-liferation using paclitaxel-loaded poly(ethylene oxide)-poly(lactide/glycolide)nanospheres. J. Biomed. Mater. Res. 42:331–38

23. Alexis F, Venkatraman SS, Rath SK, Boey F. 2004. In vitro study of releasemechanisms of paclitaxel and rapamycin from drug-incorporated biodegradablestent matrices. J. Control. Release 98:67–74


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

24. Burton PBJ, Yacoub MH, Barton PJR. 1998. Rapamycin (sirolimus) inhibitsheart cell growth in vitro. Pediatr. Cardiol. 19:468–70

25. Kahan BD, Camardo JS. 2001. Rapamycin: clinical results and future opportu-nities. Transplantation. 72:1181–93

26. Morice MC, Serruys PW, Sousa JE, Perin M, Colombo A, et al. 2002. RAVELstudy group, a randomized comparison of a sirolimus-eluting stent with a stan-dard stent for coronary revascularization. N. Engl. J. Med. 346:1773–80

27. Regar E, Laarman G, Blanchard D, Eltchaninoff H, Sousa JE, et al. 2002.Sirolimus. inhibits restenosis irrespective of the vessel size: a subanalysis of themulticenter RAVEL trial. J. Am. Coll. Cardiol. 39:58A

28. Degertekin M, Regar E, Tanabe K, Smits PC, van der Giessen WJ, et al. 2003.Sirolimus-eluting stent for treatment of complex in-stent restenosis, the firstclinical experiences. J. Am. Coll. Cardiol. 41:184–89

29. Regar E, Serruys PW, Bode C, Holubarsch C, Guermonprez JL, et al. 2002.Angiographic findings of the multicenter randomized study with the sirolimus-eluting Bx velocity balloon expandable stent (RAVEL): sirolimus-eluting stentsinhibit restenosis irrespective of the vessel size. Circulation 106:1949–56

30. Moses JW, Leon MB, Popma JJ, Fitzgerald PJ, Holmes DR, et al. 2003.Sirolimus-eluting stents versus standard stents in patients with stenosis in anative coronary artery. N. Engl. J. Med. 349:1315–23

31. Grube E, Silber S, Hauptmann KE, Mueller R, Buellesfeld L, et al. 2003.TAXUS I: six- and twelve-month results from a randomized, double-blind trialon a slow-release paclitaxel-eluting stent for de novo coronary lesions. Circula-tion 107:38–42

32. Tanabe K, Serruys PW, Degertekin M, Guagliumi G, Grube E, et al. 2004.Chronic arterial responses to polymer-controlled paclitaxel-eluting stents: com-parison with bare metal stents by serial intravascular ultrasound analysis: datafrom the randomized TAXUS-II trial. Circulation 109:196–200

33. Tanabe K, Serruys PW, Grube E, Smits PC, Selbach G, et al. 2003. TAXUS IIItrial: in-stent restenosis treated with stent-based delivery of paclitaxel incorpo-rated in slow-release polymer formulation. Circulation 107:559–64

34. Stone GW, Ellis SG, Cox DA, Hermiller J, O’Shaughnessy C, et al. 2004. Apolymer-based, paclitaxel-eluting stent in patients with coronary artery disease.N. Engl. J. Med. 350:221–31

35. Tsuji T, Tamai H, Igaki K, Kyo E, Kosuga K, et al. 2003. Biodegradable stentsas a platform to drug loading. Int. J. Cardiovasc. Interv. 5:13–16

36. Talja M, Valimaa T, Tammela T, Petas A, Tormala P. 1997. Bioabsorbable andbiodegradable stents in urology. J. Endourol. 11:391–97

37. Kapoor R, Liatsikos EN, Badlani G. 2000. Endoprostatic stents for managementof benign prostatic hyperplasia. Curr. Opin. Urol. 10:19–22

38. Kletscher BA, Oesterling JE. 1995. Prostatic stents. Current perspectives for themanagement of benign prostatic hyperplasia. Urol. Clin. North Am. 22:423–30

39. de la Rosette JJMC, Beerlage HP, Debruyne FMJ. 1997. Role of temporarystents in alternative treatment of benign prostatic hyperplasia. J. Endourol.11:467–72


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

40. Kemppainen E, Talija M, Riihela M, Pohjonen T, Tormala P, Alfthan O. 1993.A bioresorbable urethral stent. An experimental study. Urol. Res. 21:235–38

41. Petas A, Talja M, Tammela T, Taari K, Lehtoranta K, et al. 1997. A randomizedstudy to compare biodegradable self-reinforced polyglycolic acid spiral stentsto suprapubic and indwelling catheters after visual laser ablation of the prostate.J. Urol. 157:173–76

42. Petas A, Isotalo T, Talja M, Tammela TLJ, Valimaa T, Tormala P. 2000. Arandomised study to evaluate the efficacy of a biodegradable stent in the pre-vention of postoperative urinary retention after interstitial laser coagulation ofthe prostate. Scand. J. Urol. Nephrol. 34:262–66

43. Petas A, Karkkainen P, Talja M, Taari K, Laato M, et al. 1997. Effectsof biodegradable self-reinforced polyglycolic acid, poly-DL-lactic acid andstainless-steel spiral stents on uroepithelium after Nd:YAG laser irradiationof the canine prostate. Br. J. Urol. 80:903–7

44. Lumiaho J, Heino A, Pietilainen T, Ala-Opas M, Talja M, et al. 2000. Themorphological, in situ effects of a self-reinforced bioabsorbable polylactide (SR-PLA 96) ureteric stent; an experimental study. J. Urol. 164:1360–63

45. Lumiaho J, Heino A, Tunninen V, Ala-Opas M, Talja M, et al. 1999. Newbioabsorbable polylactide ureteral stent in the treatment of ureteral lesions: anexperimental study. J. Endourol. 13:107–12

46. Rafanan AL, Mehta AC. 2000. Stenting of the tracheobronchial tree. Radiol.Clin. North Am. 38:395–408

47. Robey TC, Eiselt PM, Murphy HS, Mooney DJ, Weatherly RA. 2000.Biodegradable external tracheal stents and their use in a rabbit tracheal re-construction model. Laryngoscope 110:1936–42

48. Lochbihler H, Hoelzi J, Dietz H. 1997. Tissue compatibility and biodegradationof new absorbable stents for tracheal stabilization: an experimental study. J.Pediatr. Surg. 32:717–20

49. Filler RM, Forte V, Farga JC, Matute J. 1995. The use of expandable metal-lic airway stents for tracheobronchial obstruction in children. J. Pediatr. Surg.30:1050–55

50. Furman RH, Backer CL, Dunham M, Donaldson J, Mavroudis C, Holinger LD.1999. The use of balloon-expandable metallic stents in the treatment of pedi-atric tracheomalacia and bronchomalacia. Arch. Otolaryngol. Head Neck Surg.125:203–7

51. Lamber R. 2000. Treatment of esophagogastric tumors. Endoscopy 32:322–3152. Morgan R, Adam A. 2001. Use of metallic stents and balloons in the esophagus

and gastrointestinal tract. J. Vasc. Interv. Radiol. 12:283–9753. Nicholson T. 2000. Other uses of nonvascular stents. Hosp. Med. 61:97–10254. Kimura T, Yokoi H, Nakagawa Y. 1996. Three-year follow-up after implanta-

tion of metallic coronary artery stents. N. Engl. J. Med. 334:561–6655. Eberhart RC, Su SH, Kytai TN, Zilberman M, Liping T, et al. 2003. Biore-

sorbable polymeric stents: current status and future promise. J. Biomater. Sci.Polym. Ed. 14(4):299–312

56. Nguyen K, Su SH, Zilberman M, Bohluli P, Frenkel P, et al. 2004. Biomaterials


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

and stent technology. In Tissue Engineering and Novel Delivery Systems, ed. MYaszemski, D Trantolo, KU Lewandrowski, V Hasirci, D Altobelli, D Wise,5:107–30. New York: Marcel Dekker

57. Kohn J, Langer R. 2000. Bioresorbable and bioerodible materials. In Biomateri-als Science—An Introduction to Materials in Medicine, ed. BD Ratner, AS Hoffman,FJ Schoen, JE Lemons, pp 64–73. New York: Academic

58. Leenslag JW, Pennings AJ, Bos RRM, Rosema FR, Boering G. 1987. Resorbablematerials of poly(l-lactic acid). VI. Plates and screws for internal fracture fixa-tion. Biomaterials 8:70–73

59. Bergsma EJ, Rozema FR, Bos RRM, de Bruijn WC. 1993. Foreign body reac-tions to resorbable poly(l-lactide) bone plates and screws used for the fixationof unstable zygomatic fractures. J. Oral Maxillofac. Surg. 51:666–70

60. Bergsma EJ, de Bruijn WC, Rozema FR, Bos RRM, Boering G. 1995. Latedegradation tissue response to poly(l-lactide) bone plates and screws. Biomate-rials 16:25–31

61. Agrawal CM, Haas KF, Leopold DA, Clark H. 1992. Evaluation of poly(l-lacticacid) as a material for intravascular polymeric stents. Biomaterials 13:176–82

62. Pachence JM, Kohn J. 2000. Bioresorbable polymers for tissue engineering. InPrinciples of Tissue Engineering, ed. RP Lanza, R Langer, WL Chick, pp. 267–70.San Diego: Academic

63. Lewis DH. 1990. Controlled release of bioactive agents from lactide/glycolidepolymers. In Biodegradable Polymers as Drug Delivery Systems, ed. M Chasin, RLanger, pp. 1–41. New York: Marcel Dekker

64. Leelarusamee N, Howard SA, Malango CJ, Ma JK. 1998. A method for thepreparation of polylactic acid microencapsules of controlled particle size anddrug loading. J. Microencapsul. 5:147–57

65. Gorman S, Tunney M. 1997. Assessment of encrustation behavior on urinarytract biomaterials. J. Biomater. Appl. 12:136–66

66. Petas A, Vuopio-Varkila J, Siitonen A, Valimaa T, Talja M, Taari K. 1998.Bacterial adherence to self-reinforced polylactic acid 96 urological spiral stentsin vitro. Biomaterials 19:677–81

67. Multanen M, Talja M, Hallanvuo S, Siitonen A, Valimaa T, et al. 2000. Bacterialadherence to ofloxacin-blended polylactone-coated self-reinforced l-lactic acidpolymer urological stents. BJU Int. 86:966–69

68. Zidar J, Lincoff A, Stack R. 1999. Biodegradable stents. In Textbook of Interven-tional Cardiology, ed. EJ Topol, pp. 787–802. Philadelphia: WB Saunders. 3rded.

69. Stack RS, Califf RM, Phillips HR, Pryor DB, Quigley PJ, et al. 1988. Inter-ventional cardiac catheterization at Duke Medical Center: new interventionaltechnology. Am. J. Cardiol. 2:3F–24

70. Gao R, Shi R, Qiao S, Song L, Li Y. 1996. A novel polymeric local heparindelivery stent: initial experimental study. J. Am. Coll. Cardiol. 27:85A

71. Yamawaki T, Shimokawa H, Kozai T, Miyata K, Higo T, et al. 1998. Intra-mural delivery of a specific tyrosine kinase inhibitor with biodegradable stentsuppresses the restenotic changes of the coronary artery in pigs in vivo. J. Am.Coll. Cardiol. 32:780–86


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

72. Tamai H, Igaki K, Kyo E, Kosuga K, Kawashima A, et al. 2000. Initial and6-month results of biodegradable poly-l-lactic acid coronary stents in humans.Circulation 102:399–404

73. Zilberman M, Eberhart R, Schwade N. 2002. In vitro study of drug loadedbioresorbable films and support structures. J. Biomater. Sci. Polym. Ed. 13:1221–40

74. Nuutinen JP, Valimaa T, Clerc C, Tormala P. 2002. Mechanical properties andin vitro degragation of bioresorbable knitted stents. J. Biomater. Sci. Polym. Ed.13(12):1313–23

75. Saito Y, Minami K, Kobayashi M, Nakao Y, Omiya H, et al. 2002. New tubularbioabsorbable knitted airway stent: biocompatibility and mechanical strength.J. Thorac. Cardiovasc. Surg. 123:161–67

76. Tamai H, Igaki K, Tsuji T, Miyata K, Higo T, et al. 1999. A biodegradablepoly-l-lactic acid coronary stent in porcine coronary artery. J. Interv. Cardiol.12:443–49

77. Schwartz RS, Huber KC, Murphy JG. 1992. Restenosis and the proportionalresponse to coronary artery injury: results in a porcine model. J. Am. Cardiol.19:267–74

78. Karas SP, Gravanis MB, Santoian EC. 1992. Coronary intimal proliferationafter balloon injury and stenting in swine: an animal model of restenosis. J. Am.Coll. Cardiol. 20:467–74

79. Tsuji T, Tamai H, Kyo E. 1998. The effect of PLLA stent design on neointimalhyperplasia. J. Cardiol. 32:235A

80. Tamai H, Igaki K, Tsuji T. 1999. A biodegradable poly-l-lactic acid coronarystent in procine coronary artery. J. Interv. Cardiol. 12:443–50

81. Tsuji T, Tamai H, Igaki K, Kyo E, Kosuga K, et al. 2001. Biodegradable poly-meric stents. Curr. Interven. Cardiol. Rep. 3:10–17

82. Miyazawa K, Kikuchi S, Fukuyama J. 1995. Inhibition of PDGF and TGF-beta 1- induced collagen synthesis, migration and proliferation by tranilast invascular smooth muscle cells. Atherosclerosis 118:213–21

83. Fukuyama J, Ichikawa K, Hamano S. 1996. Tranilast suppresses the vascularintimal hyperplasia after balloon injury in rabbits fed in a high cholesterol diet.Eur. J. Pharmacol. 318:327–32

84. Miyazawa N, Umemura K, Kondo K. 1997. Effects of pemirolast and tranilaston intimal thickening after arterial injury in the rat. J. Cardiovasc. Pharmacol.30:157–62

85. Ishiwata S, Verheye S, Robinson KA. 2000. Inhibition of neointima formationby tranilast in pig coronary arteries after balloon angioplasty and stent implan-tation. J. Am. Coll. Cardiol. 35:1331–37

86. Uurto I, Mikkonen J, Parkkinen J, Keski-Nisula L, Nevalainen T, et al. 2005.Drug-eluting bioresorbable poly-d/l-lactic acid vascular stents: an experimentalpilot study. J. Endovasc. Ther. 12:371–79

87. Zilberman M, Schwade ND, Eberhart RC. 2004. Protein-loaded bioresorbablefibers and expandable stents: mechanical properties and protein release. J.Biomed. Mater. Res. Part B Appl. Biomater. 69B:1–10


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

ANRV281-BE08-05 ARI 12 June 2006 17:11

88. Su S, Chao RY, Landau CL, Nelson KD, Timmons RB, et al. 2003. Expandablebioresorbable endovascular stent. I. Fabrication and properties. Ann. Biomed.Eng. 31:667–77

89. Zilberman M. 2006. Novel composite fiber structures to provide drug deliv-ery for medical implants and tissue regeneration application. Acta Biomater.Submitted

90. Levy Y, Zilberman M. 2006. Nove bioresorbable fiber structures loaded withproteins for tissue regeneration applications: microstructure and protein release.J. Biomed. Mater. Res. (Pt. A). In press

91. Zilberman M, Shraga I. 2006. Microsphere-based bioresorbable structuresloaded with Proteins for tissue regeneration applications. J. Biomed. Mater. Res.(Pt. A). In press

92. Kraitzer A. 2006. Paclitaxel loaded core/shell fiber structures for stent applications.MSc thesis. Tel Aviv Univ.

93. Woods TC, Marks AR. 2004. Drug-Eluting Stents. Annu. Rev. Med. 55:169–7894. Sousa JE, Serruys PW, Costa MA. 2003. New frontiers in cardiology: drug-

eluting stents: part II. Circulation 107:2383–8995. Drachman DE, Edelman ER, Seifert P, Groothuis AR, Bornstein DA, et al.

2000. Neointimal thickening after stent delivery of Paclitaxel: Change in com-position and arrest of growth over six months. J. Am. Coll. Cardiol. 36(7):2325–32

96. Duda SH, Poerner TC, Wiesinger B, Rundback JH, Tepe G, et al. 2003. Drug-eluting stents: potential applications for peripheral arterial occlusive disease. J.Vascul. Interven. Radiol. 14:291–301

97. Zilberman M, Nelson KD, Eberhart R. 2005. Mechanical properties andin-vitro degradation of bioresorbable fibers and expandable fiber-based stents.J. Biomed. Mater. Res. Part B Appl. Biomater. 74B:792–99

98. Greisler HP, Ellinger J, Schwarcz TH, Golan J, Raymond RM, Kim DU. 1987.Arterial regeneration over polydioxanone prostheses in the rabbit. Arch. Surg.,Chicago. 122:715–21

99. Makela EA, Vainionpaa S, Vihtonen K, Mero M, Helevirta P, et al. 1989. Theeffect of a penetrating biodegradable implant on the growth plate: an experi-mental study on growing rabbits with special reference to polydioxanone. Clin.Orthop. 241:300–8

100. Ye YW, Landau C, Willard JE, Rajasubramanian G, Moskowitz A, et al. 1998.Bioresorbable microporous stents deliver recombinant adenovirus gene transfervectors to the arterial wall. Ann. Biomed. Eng. 26:398–408

101. Vogt F, Stein A, Rettemeier G, Krott N, Hoffmann R, et al. 2004. Long-term assessment of a novel biodegradable paclitaxel-eluting coronary polylactidestent. Eur. Heart. J. 25:1330–40

102. Zilberman M, Schwade N, Meidell R, Eberhart R. 2001. Structured drug loadedbioresorbable films for support structures. J. Biomater. Sci. Polym. Ed. 12:875–92

103. Zilberman M. 2005. Dexamethasone loaded bioresorbable films used in medi-cal support devices: structure, degradation, crystallinity and drug release. ActaBiomater. 1(6):615–25


Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

Contents ARI 16 June 2006 15:26

Annual Reviewof BiomedicalEngineering

Volume 8, 2006Contents

Fluorescence Molecular ImagingVasilis Ntziachristos � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 1

Multimodality In Vivo Imaging Systems: Twice the Poweror Double the Trouble?Simon R. Cherry � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �35

Bioimpedance Tomography (Electrical Impedance Tomography)R.H. Bayford � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �63

Analysis of InflammationGeert W. Schmid-Schönbein � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � �93

Drug-Eluting Bioresorbable Stents for Various ApplicationsMeital Zilberman and Robert C. Eberhart � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 153

Glycomics Approach to Structure-Function Relationshipsof GlycosaminoglycansRam Sasisekharan, Rahul Raman, and Vikas Prabhakar � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 181

Mathematical Modeling of Tumor-Induced AngiogenesisM.A.J. Chaplain, S.R. McDougall, and A.R.A. Anderson � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 233

Mechanism and Dynamics of Cadherin AdhesionDeborah Leckband and Anil Prakasam � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 259

Microvascular Perspective of Oxygen-Carrying and -NoncarryingBlood SubstitutesMarcos Intaglietta, Pedro Cabrales, and Amy G. Tsai � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 289

PolymersomesDennis E. Discher and Fariyal Ahmed � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 323

Recent Approaches to Intracellular Delivery of Drugs and DNAand Organelle TargetingVladimir P. Torchilin � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 343

Running Interference: Prospects and Obstacles to Using SmallInterfering RNAs as Small Molecule DrugsDerek M. Dykxhoorn and Judy Lieberman � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 377

v

Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

Contents ARI 16 June 2006 15:26

Stress Protein Expression KineticsKenneth R. Diller � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 403

Electrical Forces for Microscale Cell ManipulationJoel Voldman � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 425

Biomechanical and Molecular Regulation of Bone RemodelingAlexander G. Robling, Alesha B. Castillo, and Charles H. Turner � � � � � � � � � � � � � � � � � � � � � � � 455

Biomechanical Considerations in the Design of Graft:The Homeostasis HypothesisGhassan S. Kassab and José A. Navia � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 499

Machine Learning for Detection and Diagnosis of DiseasePaul Sajda � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 537

Prognosis in Critical CareLucila Ohno-Machado, Frederic S. Resnic, and Michael E. Matheny � � � � � � � � � � � � � � � � � � � � 567

Lab on a CDMarc Madou, Jim Zoval, Guangyao Jia, Horacio Kido, Jitae Kim, and Nahui Kim � � � 601

INDEXES

Subject Index � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 629

Cumulative Index of Contributing Authors, Volumes 1–8 � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 643

Cumulative Index of Chapter Titles, Volumes 1–8 � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � � 646

ERRATA

An online log of corrections to Annual Review of Biomedical Engineering chapters (if any,1977 to the present) may be found at http://bioeng.annualreviews.org/

vi Contents

Ann

u. R

ev. B

iom

ed. E

ng. 2

006.

8:15

3-18

0. D

ownl

oade

d fr

om w

ww

.ann

ualr

evie

ws.

org

Acc

ess

prov

ided

by

Tel

Avi

v U

nive

rsity

on

05/0

3/16

. For

per

sona

l use

onl

y.

drug-eluting bioresorbable stents for various …meitalz/articles/d2.pdf · 2016. 5. 3. ·...

Documents