Once the physiochemical properties of the selected active compound(s) have been taken into account, the design of efficient systems for perivascular administration must consider the following criteria:
(1) administration mode, system localization and appropriate vascular constriction, which are related to the mechanical properties of the material (strength, flexibility, stability); and (2) drug release kinetics, which are related to the chemical properties of the polymer.
Administration – Localization
Different types of perivascular systems administered during open surgery are presented in Figure 3.
These systems should have adequate mechanical properties in order to be bent, sutured or injected and to remain on site for the desired period of time. Semi-solids, most frequently gels, present the advantages of easy handling and conformable covering of the graft [66]. Indeed, a gel should be easily manipulated by the surgeon to cover the anastomosis as well as the graft itself (Fig. 3 D and E). Additionally, appropriate gel viscosity would allow the gel to be administered or re-administered via ultrasound-guided injection or catheter micro-injection [67, 68]. On the other hand, gels may spread or move away from the site of application due to vascular tone, and in the case of coronary bypass, due to the contractile movements of the beating heart. However, anatomical constraints, e.g., muscle tissue/plane, might contribute to avoid gel migration as shown by magnetic resonance imaging [69]. Alternatively, to limit the spreading of a hydrophilic heparin-loaded polyvinyl alcohol (PVA) gel, Jones et al. enclosed it in a silicone shell to maintain the semi-liquid PVA-heparin at the perivascular site [9, 70]. Nevertheless, employing a silicone shell to limit the gel diffusion is not an appealing approach due to the non-degradable nature of the polymer.
Poloxamer (triblock copolymer: hydroxypoly(oxyethylene)/poly(oxypropylene)/ poly (oxyethylene) gels have been frequently tested in preclinical studies [79-86]. At concentrations of 20-30 %, these copolymers present the property of gelation above a critical temperature. They can be handled and injected easily at low temperatures as liquids, but they gel at body temperature [87]. Still, poloxamers are likely not appropriate for perivascular application due to their fast washout and clearance when in contact with body fluids. Indeed, in 1992, Simons et al. noticed that poloxamer F127 Pluronic®
disappeared after 1-2 h in all animals after application to the carotid artery [88]. Using magnetic
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
20 resonance imaging, the signals of poloxamers, applied subcutaneously, disappeared after 45 h, confirming the short retention time of this polymer [89].
Figure 3. Images of different perivascular administration systems, adapted from the literature: A:
Braided nitinol mesh [68]; B: Knitted nitinol mesh [69]; C: Provena® PET mesh, courtesy of Dr. Fr.
Saucy; D: Dipyridamole microspheres in PLGA-PEG-PLGA Regel® gel [70]; E: Sirolimus in PLGA-PEG-PLGA Regel® gel [64]; F: Sirolimus on PET mesh [71]; G: Sirolimus in PCL wrap [72]; H: Mithramycin in EVAc (polyethylene vinyl acetate) + PEG cuff [73]; I: Paclitaxel-loaded PLGA microneedle cuff [74]; J: Sunitinib-loaded PLGA wrap [75]. Copyright owners granted their permission for re-use.
A concept combining thermogelation and biodegradability has been proposed for an injectable gel system with improved safety and an appropriate retention time [90]. Cheung’s and other groups suggested the use of PLGA-PEG-PLGA gels, (PEG: polyethylene glycol) (Regel®) [67, 69, 91-93],
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
21 which are thermoresponsive and degrade within approximately 30 days [94]. This hydrophobic polymer solidifies at room temperature in contact with water within seconds upon body contact [93]
and thus requires delicate coordination in the operating room to ensure the application of its liquid form that would allow conformable covering of the graft. Another thermogelling and biodegradable gel is Atrigel®, consisting of a solution of PLGA in N-methyl-2-pyrrolidone (a class II solvent) [95].
Once injected in the body, water displaces the solvent, and the polymer precipitates instantly, forming a porous solid structure [96]. Tissue exposure to such solvents may cause local toxicity and teratogenicity [97-99]. Therefore, using a formulation with this solvent for prevention of IH is questionable.
Hyaluronic acid (HA) gels have also been used for perivascular delivery [66, 78]. HA is a natural polymer, known for its low toxicity, as it is naturally distributed throughout the connective, epithelial, and neural tissues. In its crosslinked form, mucoadhesive HA gel can avoid the migration phenomena. Its cross-linked nature slows degradation, and consequently the gel remains in vivo for several months. Importantly, cross-linked HA will eventually be enzymatically degraded by hyaluronidases. It is used for rheumatology (Ostenil®), ophthalmology (Anikavisc™) and aesthetic surgery (Fortélis Extra®) applications. Furthermore, the role of HA on IH has been described in the literature. Most studies agree that HA is able to reduce IH formation [100, 101] or at least had no effect [102]. However, an increased cellular response has been reported in some cases [103]. Two recent reviews conclude that, at high concentrations, HA is a promising candidate for IH inhibition [104, 105].
Despite the ease of administration of semi-solid forms, solid forms are frequently selected for perivascular application. Their placement is only possible when the vasculature is readily exposed during surgery, as in the case of a bypass intervention or an arteriovenous fistula placement.
Compared to semi-solid systems, solid forms are less prone to migrate away from the desired location. Non-tubular devices (cuffs and wraps) can be adapted to the size of the vessel and to angular shapes such as anastomoses. They may be sutured in place or not. The non-sutured approach allows free movement of the vessel. The sutured systems allow the surgeons to decide the constriction to apply to the vein [106]. The application of tubular devices (meshes and sheaths) is more complex and requires special equipment.
To ensure the mechanical flexibility of solid perivascular systems, the optimal thickness of PCL, PLA and PLGA wraps was set to 50 µm [75]. Polymer selection is a crucial step. Due to their high elastic modulus, cuffs and wraps made of PLGA [75, 78, 107], PLA (polylactic acid) [75] or PCL [75, 108] are often hard and brittle to handle and to suture in place, and their stiffness further increases upon drug addition [78]. For PLGA, even when implantation was possible, the wrap turned brittle after a week in vivo [75, 78]. This could result in a non-conformable covering of the graft and
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
22 could also trigger injury to the vasa vasorum, which may in turn, induce IH formation [109]. To improve elasticity, PEG or methylated PEG was added as a plasticizer to improve the elastic properties of cuffs and wraps [11, 76, 110, 111]. EVAc, usually used at 40 % vinyl acetate co-polymer composition, can form flexible matrices with satisfactory elastic properties [11, 76, 106, 112-115]. Alternatively, highly elastic poly(L-lactide-co-caprolactone) (PLCL) was used [116, 117].
This polymer has additionally an excellent elongation at break performance, allowing shape adaptation (1,000 times better than PLGA).
Overall, gels developed for perivascular administration present the significant advantage of ease of administration. Thermoresponsive or cross-linked gels are the most appealing approaches to date, ensuring injectability, low toxicity and increased retention time. While solid forms require more effort for their placement, they are undoubtedly less prone to migrate compared to gels. However, they do not envelop anastomoses, which are the sites where IH is most abundantly developed. For solid formulations, PLCL has the most interesting mechanical properties for the appropriate placement of cuffs and wraps. However, as will be discussed in section 5.3, controlling the drug release is compromised with the use of this polymer.
Vessel constriction: Stress or support?
A venous graft is subjected to shear stress and expands to adapt to the new environment of arterial pressure [118]. Following implantation, the vein graft will arterialize, meaning that wall thickening will occur to adjust to the arterial pressure [23]. Mechanical support and even constriction of the vessel to increase shear stress has been suggested as beneficial for the vessel’s patency [119]. Gels or loose-fitting wraps offer no mechanical support to the vessel. Instead, the caliber of meshes or sheaths can be selected to constrict the graft (Fig. 3 A, B, C). Even though they are technically challenging and there is no possibility to quantitate the constriction force applied, cuffs could allow the surgeon to adjust the constriction while suturing.
It is very important to carefully select the diameter of the device. However, the influence of vessel constriction on IH reduction is still a matter of debate. This issue has been investigated and reviewed by different groups comparing the effect of tight or loose-fitting MDs [23, 24, 120]. More specifically, the term tight fitting is used for devices with a diameter equal or smaller than the dilated vessel, whereas loose fitting is characterized by a lack of intimate contact with the graft. Table 2 summarizes the results from studies on the effect of constriction on IH. It was suggested that non-constrictive MDs are preferred for leaving space for the development of neoadventitial vasa vasorum [120-125]. However, it was shown that a tight-fitting MD would prevent the vein from over-distension, thus favoring lumen size matching between the graft and the host artery. A tight fit would thus increase the shear stress and reduce IH [72, 126-129]. Importantly, constriction should be
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
23 adequate to eliminate luminal irregularities in order to inhibit IH, but should not provoke vessel wall folding [71, 130].
Another important parameter regarding constricting systems is the duration of the constrictive effect.
It has been suggested that since IH is initiated within the first month following the intervention, the presence of a supportive material is not necessary after that period [120, 121, 125]. Bioresorbable systems, either semisolids or solids, would be preferred as the risk of infection linked to the application of a solid, permanent foreign body is high with potentially dramatic consequences for the patient. The nature of the biomaterial employed and particularly its biodegradability will determine: i) the duration of the constriction and ii) the retention time of the system.
When non-degradable nitinol metal meshes are used, they offer higher kink resistance and less fraying compared to polymeric meshes [72]. The choice of the filaments/wire arrangement (woven, knitted, braided) affects the radial pulse compliance, flexibility deployment and bending stress [131].
Knitted meshes have a higher pulse compliance, lower bending stiffness and an overall greater capacity to maintain a constant caliber diameter [72]. It was also shown that the type of weaving of the meshes impacts the in vivo effect. Indeed, the remaining medial muscle mass was superior in knitted compared to braided meshes [72].
In contrast, if biodegradable polymers are used, the constraint effect will evolve with time [22]. This parameter might be appealing, since it has been suggested that ideally the transfer of shear stress load to the vessel should occur gradually [132]. Indeed, biodegradable materials may cease to fulfill their supportive role within the first month following application due to alterations in the polymer crystallinity, making the device stiffer as biodegradation occurs. Researchers observed that PLGA wraps were stiff upon explantation after 14-28 d [75, 78]. If a stiff material remains perivascularly for such a long period of time, concerns are raised over its mechanical properties and over the impact on the prevention of IH. It has also been observed that polymer hardening generated thrombosis [75].
The data presented in Table 2 suggest that constriction is beneficial against IH. Several groups have shown that IH can be sufficiently reduced only with the use of devices offering mechanical support [71, 72, 117, 120-125, 128-130, 133-138]. It seems, however, that the biodegradable materials currently available will not fulfill the constriction demands of the perivascular systems. New materials need to be explored.
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
24 Table 2. Effects of solid perivascular devices (mesh and sheath) on IH after an arteriovenous bypass intervention. The in vivo effect of intimal hyperplasia reduction is reported here as the percent reduction in intima thickness. P < 0.05 (*), P < 0.01 (**), P < 0.001 (***). Refer to the source for statistical analysis method. Qc = quotient of cross-sectional area of host artery to vein graft.
External support Animal model In vivo effect on IH Ref
Non-constrictive
PTFE sheath (vascular graft) Saph. vein interposition on carotid, pig -72 % (***) [119]
PET mesh VS PTFE sheaths Saph. vein interposition on carotid, pig IH reduction for PET [130]
PET mesh Saph. vein interposition on carotid, pig -72 % (**) 8 mm mesh [120]
PET mesh Saph. vein interposition on carotid, pig -97 % (**) [121]
PET mesh Saph. vein interposition on carotids, pig -42 % (*) [118]
PLGA sheath Saph. vein interposition on carotids, pig -73 % (*) [122]
Constrictive
Polypropylene mesh Jug. vein interposition on carotid, dog IH reduction [133]
PTFE sheath (vascular graft) Jug. vein interposition on carotid, rabbit IH reduction [132]
PTFE sheath (vascular graft) Saph. vein interposition on carotid, pig No IH reduction [133]
PLA sheath Saph. vein interposition on carotids, pig No IH reduction [134]
Polyurethane/PLA sheath Jug. vein interposition on carotid, rabbit IH reduction [135]
Collagen sheath Jug. vein interposition on carotid, rabbit -45 % (***) [125]
Knitted Nitinol mesh Femorofemoral interposition, baboon -88 % (***) [69]
Knitted Nitinol mesh Aortocoronary bypass, baboon -60 % (***) Qc=0.41 [136]
Braided alloy mesh Aortocoronary bypass, sheep -52 % (*) [137]
Braided Nitinol mesh Femorofemoral interposition, baboon -91 % (**) Qc=1.47 or 3.09 [68]
Comparison
PTFE sheath (vascular graft) Jug. vein interposition on carotid, rabbit IH reduction for tight fit [126]
Braided Nitinol mesh Femorofemoral interposition, baboon -98 % (**) Qc=0.45 or 1.16 [138]
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
25 Drug release profile from the DDSs
The need for a controlled drug release
Although drug-free MDs such as meshes have proven their efficacy in controlling focal vascular pathologies, their efficiency is reinforced when active compounds are added to the device. Skalsky et al. compared the effect of a bare polyester mesh and a sirolimus-releasing mesh wrapped around an autologous vein grafted in a rabbit model. Although three weeks post-surgery, no difference was seen between both devices, after six weeks, only the sirolimus-eluting meshes ensured the desired effect [74]. A wide spectrum of pharmaceutical formulations has been designed for the perivascular application of bioactive compounds. The choice of the compound, which is of utmost importance, has been reviewed elsewhere [15-17]. In this section, we focus on the importance to deliver the compound at the right place with the right kinetics.
The use of biodegradable materials offers the possibility of controlling the drug release rate. Most of them were initially investigated for perivascular application to inhibit thrombosis following vascular surgery [70, 142]. Heparin was administered perivascularly, from fast-releasing hydrophobic polyanhydride wraps. More than a two-fold improvement in anastomotic patency was observed after 7 days, compared to the untreated control in a rat microvascular thrombosis model [142]. Rogers et al. reported that the sustained perivascular release of heparin from an EVAc matrix had not only an effect on thrombosis reduction but also reduced intimal thickness by 50 % in a rabbit’s iliac artery injury model [143, 144].
Apart from these early studies, perivascular formulations are mainly used to inhibit IH, which is described as a three-stage process (Fig. 1C). The most critical processes of the IH development occur in the first 4 weeks following the vascular injury. Ideally, a single application of a DDS should be sufficient to prevent IH. For this, the drug’s availability has to be appropriate to fit the time course of the evolution of the pathology [56, 145, 146]. In the literature describing perivascular DDS, the duration of the drug release ranges from a few hours to 3 months. Most of them focus on a 2- to 4-week duration. The importance of sustaining perivascular release was demonstrated in rats, as the fast uncontrolled release of high doses of heparin led to lethal bleeding as opposed to a controlled release formulation that allowed inhibition [147].
The selection of the drug release profile (e.g., diffusion, zero-order, biphasic) is equally as important as the release duration. A diffusion-like release should ensure that the burst is controlled, to avoid exposing the surrounding tissues to excessive or even toxic levels of drug. A release profile with a lag phase with no drug release might not be desirable for this pathology. We recently documented that the combination of a burst and a sustained release are necessary for atorvastatin to effectively reduce IH by 68 % [66]. By combining atorvastatin’s diffusion-driven release from a gel with PLGA
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
26 microparticles release profile, the lag phase was avoided, and the treatment was efficient. Park has also noted the importance of an adequate drug release profile in perivascular delivery [148]. Table 3 summarizes the DDS found in the literature. Gels (with or without particulate systems) and solid formulations have been largely studied for drug release for perivascular application. This table investigates the relationship between the drug release duration in vitro and the in vivo efficacy of the drug for IH reduction.
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
27 Table 3: Summary of DDS found in the literature for perivascular administration. In vitro release duration is indicated when available. The in vivo effect is presented as IH reduction as I/M: intima/media ratio or I: Intima reduction or L: Lumen increase compared to control. P < 0.05 (*), P < 0.01 (**), P <
0.001(***). Refer to the source for the statistical analysis method. PECC: poly(ethylene carbonate-ε-caprolactone), PEAD-SA: poly-(erucic acid dimer-sebacic anhydride), NA: data not available, PPO: Polypropylene Oxide.
Drug Delivery System Active compound In Vitro Release In Vivo Effect Ref Gels
Poloxamer Difluoromethylornithine NA L +49 % [77]
Poloxamer Suramin NA I -50 % (*) [80]
Poloxamer Cilostazol NA I/M -62 % (*) [81]
Poloxamer Prolif. cell nuclear antigen NA I/M -27 % (*) [82]
Poloxamer CDK2 oligonucleotides NA I/M -55 % (*) [146]
Poloxamer c-myb oligonucleotides NA I/M -63 % [85, 147]
Poloxamer c-myb oligonucleotides 3 d (diffusion) I/M -90 % [143]
Poloxamer c-myc oligonucleotides 3 d (diffusion) I/M +30 % [143]
Poloxamer Nitric Oxide donor 70-80 d I/M -46 % (*) [148, 149]
Poloxamer Sirolimus NA I/M -55 % (**) [78]
Poloxamer Sirolimus NA I -41 % (ns) [79]
Cross-linked HA Atorvastatin 3 d (diffusion) I/M -30 % (ns) [63]
Regel® Sirolimus 14 d (zero-order) L +52 %(*) [64, 66]
Regel® Paclitaxel NA I -47 % (*) [88]
Atrigel® Nitric oxide donor 14 d I/M -75 % (*) [150]
Atrigel® Nitric oxide donor t1⁄2: 39 min I -39 % (**) [93]
Fibrin glue Losartan NA I -82 % (*) [151]
PVAL in silastic shell Heparin NA I/M -51 % (**) [6]
PEG-Cys-NO photopolymerized Nitric Oxide donor NA I/M -77 % (*) [152, 153]
Peptide amphiphile nanofiber Nitric Oxide donor 5 d (diffusion) I/M -77 % (*) [154]
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
28
Gels + particles
PLGA nanoparticles in Poloxamer Sirolimus 30 d (biphasic) I/M -45 %(*) [76]
PLGA microparticles in Regel® Dipyridamole 35 d (zero-order) not effective [70, 90]
PLGA microparticles in clHA gel Atorvastatin in micr. 45 d (biphasic) I/M -39 % (ns) [63]
PLGA microparticles in clHA gel Atorvastatin in gel/micr. 45 d (diffusion) I/M -68 % (*) [63]
PLGA microparticles in Alginate gel Heparin 36 d (diffusion) I/M -62 % (**) [144]
PVAL microparticles in Poloxamer Sirolimus 9 d (zero-order) I 1 w -50 % (*), 4 w:(ns) [155]
PLGA/polyeth. oxide nanoparticles Paclitaxel 28 d (diffusion) NA [156]
Calcium alginate microparticles Fibroblast growth factor 1-2 h NA [26]
Solids
EVAc matrix Heparin 12 d (diffusion) L +62 % (*) [157]
EVAc matrix c-myb oligonucleotides 15 d (zero-order) I/M -72 % (*) [143]
EVAc matrix c-myc oligonucleotides 15 d (zero-order) I/M -94 % (*) [143]
EVAc matrix Heparin 16 d (diffusion) I -55 % (**) [141]
EVAc cuff AG-17 tyrosine kinase 21 d (in vivo) I/M -57 % (*) [109]
EVAc cuff Verapamil <24 d (zero-order) I/M -31 % (*) [110]
EVAc cuff Colchicine 30 d (diffusion) I/M -73 % (*) [103]
EVAc wrap Paclitaxel 30 d (diffusion) L +15 % (*) [112]
EVAc + PEG cuff Mithramycin 30 d (in vivo) I/M -48 % (*) [8]
EVAc + PEG wrap Paclitaxel 20 d (diffusion) L +100 % (ns) [73, 158]
PLGA cuff Paclitaxel <7 d I -91 % (*) [104]
PLGA wrap Sunitinib 40 d (biphasic) NA [75]
PLGA wrap Rapamycin 50 d (biphasic) thrombosis [72]
PLGA sheath + fibrin glue Bosentan - I/M -71 % [159]
PLGA microneedle cuff Paclitaxel 7 d (diffusion) I/M -53 % (*) [74, 160]
PLGA + PDLLA-MePEG cuff Paclitaxel 30 d (diffusion) NA [107]
PLLA wrap Rapamycin >50 d (zero-order) Thrombosis [72]
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
29
PCL wrap Sirolimus 50 d (zero-order) I/M -85 %(*) [72]
PCL+PEG wrap Paclitaxel >21 d (zero-order) I -76 % (**) [105]
PCL+PEG wrap Sirolimus >21 d (zero-order) I -75 % (**) [105]
PCL+PEG wrap Blebbistatin 21 d (ex vivo) I -73 %(*) [108]
PCL+mPEG sheath Paclitaxel 30 d (diffusion) L -2 % (ns) [112]
PCL+mPEG paste Paclitaxel 30 d (diffusion) L +20 % (*) [112]
PLCL wrap Sirolimus not clear I/vessel -39 % (*) [113]
PLCL on polyester mesh Sirolimus 42 d (diffusion) I/M -70 % (**) [71, 114]
PLGA-PPO-PLGA in PLCL nanofib. Paclitaxel 90 d NA [161]
PECC nanofibers Paclitaxel >3 m (diffusion) NA [162]
Fibrin glue losartan >4 h I/M -82 %(*) [151]
poly(diol-co-citrate) cuff Nitric Oxide donor 3 d (zero-order) I/M -38 % (*) [163, 164]
poly(diol-co-citrate) wrap all-trans retinoic acid 14 d (diffusion) I/M -57 %(*) [165]
PES matrix with PLA coat 2 % Heparin 20 d (zero-order) I/M -54 % (*) [166]
cross-linked PEG wrap Cyclosporin A 50 h (diffusion) NA [167]
Silicone cuff dexamethasone 92 d (diffusion) I/M -75 % (***) [168]
CHAPTER I - Perivascular medical devices and drug eluting systems: Making the right choices
30 Controlled release systems
A wide range of controlled release systems has been described in the literature. There are biodegradable polymeric systems such as gels, microparticles and wraps encapsulating drugs.
Formulations with a drug dispersed in a poloxamer or HA gels do not sustain the release for more than 3 days [81-85]. Cross-linked HA hydrogel mixed with atorvastatin gave a fast release over 3 days in vitro and reduced IH only by 24 % (intima/media) [66]. Similarly, the use of poloxamer did not provide a sufficiently sustained release of c-myc for IH suppression (-30 %) [146]. Regel® and Atrigel® are conceived to sustain the drug release, thanks to the presence of PLGA in the formulation [67, 79, 91, 92, 153]. Despite the fact that in vitro sirolimus release from Regel® was sustained over 14 days, weekly replenishments were necessary in vivo for the formulation to be efficient [69].
Formulations with a drug dispersed in a poloxamer or HA gels do not sustain the release for more than 3 days [81-85]. Cross-linked HA hydrogel mixed with atorvastatin gave a fast release over 3 days in vitro and reduced IH only by 24 % (intima/media) [66]. Similarly, the use of poloxamer did not provide a sufficiently sustained release of c-myc for IH suppression (-30 %) [146]. Regel® and Atrigel® are conceived to sustain the drug release, thanks to the presence of PLGA in the formulation [67, 79, 91, 92, 153]. Despite the fact that in vitro sirolimus release from Regel® was sustained over 14 days, weekly replenishments were necessary in vivo for the formulation to be efficient [69].