Scholarly article on topic 'Drug and cell delivery for cardiac regeneration'

Drug and cell delivery for cardiac regeneration Academic research paper on "Basic medicine"

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{"Myocardial infarction" / "Heart failure" / "Cell therapy" / "Growth factor" / Biomaterials / "Medical device" / "Drug delivery" / "Regenerative medicine"}

Abstract of research paper on Basic medicine, author of scientific article — Conn L. Hastings, Ellen T. Roche, Eduardo Ruiz-Hernandez, Katja Schenke-Layland, Conor J. Walsh, et al.

Abstract The spectrum of ischaemic cardiomyopathy, encompassing acute myocardial infarction to congestive heart failure is a significant clinical issue in the modern era. This group of diseases is an enormous source of morbidity and mortality and underlies significant healthcare costs worldwide. Cardiac regenerative therapy, whereby pro-regenerative cells, drugs or growth factors are administered to damaged and ischaemic myocardium has demonstrated significant potential, especially preclinically. While some of these strategies have demonstrated a measure of success in clinical trials, tangible clinical translation has been slow. To date, the majority of clinical studies and a significant number of preclinical studies have utilised relatively simple delivery methods for regenerative therapeutics, such as simple systemic administration or local injection in saline carrier vehicles. Here, we review cardiac regenerative strategies with a particular focus on advanced delivery concepts as a potential means to enhance treatment efficacy and tolerability and ultimately, clinical translation. These include (i) delivery of therapeutic agents in biomaterial carriers, (ii) nanoparticulate encapsulation, (iii) multimodal therapeutic strategies and (iv) localised, minimally invasive delivery via percutaneous transcatheter systems.

Academic research paper on topic "Drug and cell delivery for cardiac regeneration"

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ADR-12647; No of Pages 22

Advanced Drug Delivery Reviews xxx (2014) xxx-xxx

Advanced DRUG DELIVERY

Reviews

i Drug and cell delivery for cardiac regeneration^

qi Conn L. Hastings a,b,c,1 Ellen T. Roche d,e,1 Eduardo Ruiz-Hernandez a,b,c, Katja Schenke-Laylandf,g,h,

3 Conor J. Walsh d,e, Garry P. Duffy aÄc,*

4 a Tissue Engineering Research Group, Dept. of Anatomy, Royal College of Surgeons in Ireland (RCSI), 123 St. Stephens Green, Dublin2, Ireland

5 b Trinity Centre for Bioengineering, Trinity College Dublin (TCD), College Green, Dublin 2, Ireland

6 c Advanced Materials and Bioengineering Research (AMBER) Centre, RCSI & TCD, Dublin 2, Ireland

7 d School of Engineering and Applied Sciences, Harvard University, 29 Oxford Street, Cambridge, MA 02138, USA

8 e Wyss Institute for Biologically Inspired Engineering, 60 Oxford Street, Cambridge, MA 02138, USA

9 f Dept. of Women's Health, University Women's Hospital, Eberhard-Karls-University Tübingen, 72076 Tübingen, Germany

10 g Dept. of Cell and Tissue Engineering, Fraunhofer Institute for Interfacial Engineering and Biotechnology (IGB), 70569 Stuttgart, Germany

11 h Dept. of Medicine/Cardiology, Cardiovascular Research Laboratories, David Geffen School of Medicine at UCLA, Los Angeles, CA, USA

12 ARTICLE INFO

ABSTRACT

20 21 22 23

14 Available online xxxx

Keywords:

Myocardial infarction Heart failure Cell therapy Growth factor Biomaterials Medical device Drug delivery Regenerative medicine

The spectrum of ischaemic cardiomyopathy, encompassing acute myocardial infarction to congestive heart 24 failure is a significant clinical issue in the modern era. This group of diseases is an enormous source of morbidity 25 and mortality and underlies significant healthcare costs worldwide. Cardiac regenerative therapy, whereby pro- 26 regenerative cells, drugs or growth factors are administered to damaged and ischaemic myocardium has demon- 27 strated significant potential, especially preclinically. While some of these strategies have demonstrated a measure 28 of success in clinical trials, tangible clinical translation has been slow. To date, the majority of clinical studies and a 29 significant number of preclinical studies have utilised relatively simple delivery methods for regenerative thera- 30 peutics, such as simple systemic administration or local injection in saline carrier vehicles. Here, we review cardiac 31 regenerative strategies with a particular focus on advanced delivery concepts as a potential means to enhance 32 treatment efficacy and tolerability and ultimately, clinical translation. These include (i) delivery of therapeutic 33 agents in biomaterial carriers, (ii) nanoparticulate encapsulation, (iii) multimodal therapeutic strategies and (iv) 34 localised, minimally invasive delivery via percutaneous transcatheter systems. 35

© 2014 Published by Elsevier B.V.

Contents

43 1. Introduction..............................................................................................................................0

44 2. Cell therapy..............................................................................................................................0

45 2.1. Introduction to cardiac cell therapy..................................................................................................0

46 2.1.1. Bone marrow derived stem cells — heterogeneous populations (BMMNCs)......................................................0

47 2.1.2. Purified stem cell populations: MSCs and EPCs................................................................................0

48 2.1.3. Skeletal myoblasts........................................................................................................0

49 2.1.4. Cardiac stem cells..........................................................................................................0

50 2.1.5. Cardiopoietic stem cells....................................................................................................0

51 2.2. Additional considerations for cell therapy............................................................................................0

52 2.3. Cells with biomaterial carriers......................................................................................................0

Abbreviations: MI, myocardial infarction; CHF, congestive heart failure; CSCs, cardiac stem cells; BMMNCs, bone marrow derived mononuclear cells; MSCs, human mesenchymal stem cells; VEGF, vascular endothelial growth factor; GCSF, granulocyte colony stimulating factor; ADSCs, adipose derived stem cells; HGF, hepatocyte growth factor; (-GP, (-glycerophosphate; HEC, hydroxy-ethyl cellulose; PEG, poly(ethylene glycol); PCL, polycaprolactone; ECM, extracellular matrix; RGD, Arg-Gly-Asp; PG, poly(e-caprolactone)/gelatin; LVEDVI, left ventricular end diastolic volume index; LVR left ventricular restraint; BCM, bioabsorbable cardiac matrix; PGE2, prostaglandin E2; PGI2, prostaglandin I2; PLGA, polylactic-co-glycolic acid; PP, pyrvinium pamoate; NADH, nicotinamide adenine dinucleotide; DPP-IV, dipeptidylpeptidase IV; miR microRNA; modRNA, modified RNA; LVEF, left ventricular ejection fraction; PEI, polyethylenimine; APOSEC, apoptotic peripheral blood cells; SDF-1, stromal cell derived factor-1; NRG-1, neuregulin-1; IGF-1, insulin-like growth factor-1; FGF-1, fibroblast growth factor-1; Shh, Sonic hedgehog morphogen; Ang-1, angiopoietin-1; CPC, cardiac progenitor cell; CDC, cardiosphere derived cell; VRD, ventricular restraint device; PTCA, percutaneous trans-luminal coronary angioplasty.

☆ This review is part of the Advanced Drug Delivery Reviews theme issue on "Scaffolds, Cells, Biologics: At the Crossroads of Musculoskeletal Repair".

* Corresponding author at: Dept. of Anatomy, Royal College of Surgeons in Ireland (RCSI), 123 St. Stephens Green, Dublin 2, Ireland.

Q2 1 Authors contributed equally.

http://dx.doi.org/m1016/j.addr.2014.08.006 0169-409X/© 2014 Published by Elsevier B.V.

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2.3.1. Injectable hydrogels.................................................... 0

2.3.2. Preformed porous scaffolds................................................. 0

3. Cell-free approaches............................................................ 0

3.1. Acellular material-based scaffolds.................................................. 0

3.2. Endogenous targeting....................................................... 0

3.2.1. Small molecules...................................................... 0

3.2.2. RNA therapeutic strategies................................................. 0

3.2.3. Direct reprogramming................................................... 0

3.2.4. Growth factors and proteins................................................ 0

4. The case for advanced delivery....................................................... 0

4.1. Multimodal therapeutic strategies.................................................. 0

4.2. Minimally invasive therapy — catheter delivery............................................ 0

4.2.1. Catheters for material based approaches........................................... 0

4.3. Conclusion............................................................ 0

Acknowledgements.............................................................. 0

References.................................................................. 0

1. Introduction

This review encompasses drug and cell delivery for cardiac regeneration. This treatment can be cardioprotective; to protect heart muscle tissue after an acute myocardial infarction (MI), or cardiorestorative; to regenerate tissue in patients with chronic ischaemic heart failure. Acute myocardial infarction occurs upon occlusion of one of the coronary vessels, most commonly due to atherosclerotic plaque, resulting in an ischaemic region of myocardium which, even if reperfused, can produce lasting tissue damage with associated symptoms. Initially, Ml produces an inflammatory response and extensive ischaemic death of cardiomyocytes within the affected area, resulting in a partial loss of ventricular function. Over time, especially if the affected area is expansive and transmural, complex alterations occur in the myocardium, a phenomenon known as ventricular remodelling [1]. These adaptations are an attempt to compensate for ventricular malfunction. However, the heart possesses only a limited regenerative capacity. Remodelling encompasses the creation of collagenous, non-contractile scar tissue, thinning of the myocardial wall and progressive enlargement and

dilation of the ventricle. This ultimately contributes to a decrease in ventricular contractile function and output. This can progress to congestive heart failure (CHF), where the heart is unable to pump enough blood to meet the metabolic demands of the body [2-4].

MI represents an enormous source of morbidity and mortality on a global scale. Coronary artery diseases such as MI and CHF are the main cause of death in developed countries, and pose a substantial healthcare burden [3]. According to the European Society of Cardiology one in six men and one in seven women in Europe will die from myocardial infarction [5]. The American Heart Association reports that 635,000 Americans have a new myocardial infarction each year and that the number of deaths attributable to heart failure in the US in 2009 was 275,000 [6]. Current therapies for the treatment of MI and CHF include pharmacological intervention, surgical procedures such as ventricular resection, coronary artery bypass or mechanical aids such as left ventricular assist devices. Such approaches serve to restore function or limit disease progression to some degree, but are not always effective long-term [7]. Reperfusion of the culprit artery (with coronary angioplasty and/or stent placement) can have a profound effect on limiting infarct

cardiac cell therapy

Fig. 1. Clinical trials in cell therapy: This figure shows the range and progression of cardiac cell therapy trials, with cell type underneath (graphically represented above) and depicts the trend of moving from unselected cell populations and different cell types towards cardiopoietic and cardiac stem cells.

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size and increasing patient survival [8]. This technique can also limit ventricular remodelling with the objective of improving ventricular function and clinical outcomes. However, myocardial necrosis begins rapidly following coronary occlusion, usually before reperfusion can be accomplished [9]. Post-infarction remodelling and the progression to heart failure therefore remain a challenge in the treatment of cardiovascular disease. The most effective treatment for end-stage CHF is heart transplantation, which is limited by the availability of heart donors and also requires a highly invasive and complex surgical procedure [2,7].

This review covers cell and drug delivery, and additional cell-free approaches that share a common goal of enabling cardiac regeneration, and attenuation or prevention of negative compensatory remodelling (limiting infarct size, reducing or preventing infarct expansion and reducing ventricular wall stress). These approaches have shown promise in addressing shortcomings in conventional cardioprotective and cardiorestorative treatments for MI and CHF, respectively. However, clinical translation of regenerative therapeutics has been slow to date. Here, we suggest a perspective on how advanced delivery strategies could be synergistically engaged in the facilitation of cardiac regeneration, for enhanced efficacy and treatment tolerability, with greater potential for clinical translation.

2. Cell therapy

2.1. introduction to cardiac cell therapy

Multiple trials have been initiated addressing the transplantation of stem cell populations for cardiac regeneration. An appropriate regenerative cell population selection is critical for effective therapy. Extensive preclinical and clinical trials have investigated a number of cell types for cardiac regeneration including skeletal myoblasts, mesenchymal stem cells (bone marrow derived and adipose derived), embryonic stem cells, and cardiac stem cells. Although most cell types have produced promising results in vitro and in preclinical studies [10-21], and have been shown to be safe in clinical trials, cardiac stem cells, or cardio-poietic stem cells have shown the most promise in terms of efficacy. Thus, the trend is towards delivery of cells derived from the heart, or lineage-specified for optimal therapy for the diseased tissue. The trials are summarised in Fig. 1, and trials for each cell-type are described in the following sections.

2.1.1. Bone marrow derived stem cells — heterogeneous populations (BMMNCs)

Bone marrow aspirate or lineage-unselected bone marrow derived mononuclear cells (BMMNCs) have been used for a significant number of preliminary clinical studies. These studies have consistently demonstrated the safety and feasibility of BMMNC administration, encouraging further investigation, but clinical benefits to date have not been convincing. Orlic et al. demonstrated that intramyocardial injection of BMMNCs improved cardiac contractility and resulted in the formation of new cardiac tissue in a mouse model of MI [10,11]. Kudo et al. reported that BMMNCs could reduce infarct size and fibrosis, and differentiate into cardiomyocytes and endothelial cells [12]. However more recent research showed that these cells likely do not differentiate into cardiomyocytes [22]. Clinical trials such as TOPCARE-AMI [23], REPAIR-AMI [24], BOOST [8,25] and FINCELL [26] have shown increases in left ventricular ejection fraction (LVEF) in cell treated patients compared to controls at time points up to 18 months. Long-term (5-year) benefits were demonstrated in the TOPCARE-AMI trial [27] but not in the BOOST trial [28]. In contrast, the ASTAMI [29], BONAMI [30], LeuvenAMI [31], and HEBE [32] trials showed no significant increase in left ventricular ejection fraction over the control group. A Phase I trial (NCT00114452) [33] with prochymal allogeneic stem cells (Osiris Therapeutics Inc.) showed an increase in LVEF at 6 months after allogeneic BMMNC transplantation, but no improvement in patient physical

performance, as measured by the six minute walk test, highlighting the need for a consensus on standardized accepted metrics for cardiac cell therapy efficacy. Trials carried out by the Cardiovascular Cell Therapy Research Network (CCTRN) indicated no clinical benefit of BMMNCs in acute myocardial infarction (AMI), where they looked at timing of postAMI intracoronary administration in the TIME [34] and LateTIME [35] trials. Numerous multicentre studies are ongoing to investigate autolo-gous bone marrow cell therapy including REVITALIZE (NCT00874354), REGEN-AMI (NCT00765453), REPAIR-ACS (NCT00711542), SWISS-AMI (NCT00355186) and BAMI (NCT01569178). Similarly, no clinical benefit was noted in a trial investigating transendocardial delivery of BMMNCs for heart failure (FOCUS-CCTRN) [36], although TOPCARE-CHD [37] showed a 2.9% increase in LVEF over base-line at 3 months. The overall negative results of these trials have encouraged exploration of other cell types or "next-generation" cell therapy, where cells are subjected to screening assays to predict regenerative potential before cell transplantation [38], or cells are modified or delivered concomitantly with drugs, as will be discussed in subsequent sections. The prevailing concept of BMMNC efficiency is explained by the paracrine hypothesis, where soluble factors (chemokines, growth factors, etc.) are secreted by transplanted cells, especially in hypoxic environments, and encourage cardiac repair [39]. This hypothesis has been supported experimentally through demonstration that conditioned media can somewhat replicate the effects of stem cell therapy [40]. Potential mechanisms include increasing angio-genesis, protecting endogenous cells, attuning the inflammatory processes and encouraging cell-cycle re-entry [41].

2.1.2. Purified stem cell populations: MSCs and EPCs

More recently, bone marrow aspirate has been purified by pheno-typic features into two multipotentcell populations; human mesenchymal stem cells (hMSCs) and endothelial progenitor cells (EPCs). Purified sub-populations were demonstrated to show higher engraftment, and can induce endogenous cardiomyogenesis [42]. BMMNCs have been delivered via intracoronary injections for the treatment of acute MI, but these purified subpopulations can be used for the treatment of chronic ischaemia and refractory angina. Clinical trials have been initiated for both subpopulations. The POSEIDON trial compared autologous and allogeneic hMSC transplantation in patients with ischaemic cardio-myopathy at different doses, and showed that allogeneic cells did not elicit donor-specific immune reactions, and that both groups favourably affect patient functional capacity and ventricular remodelling, although they did not increase ejection fraction [43]. The TAC-HFT trial compared BMMNCs and hMSCs for heart failure, and reported that both were safe, with a trend towards reverse remodelling and regional contractility. Adipose tissue is also being used as a source for hMSCs. When adipose stem cells and bone marrow stem cells were compared in a porcine MI model, they both showed similar improvements in cardiac function and increased capillaries in the infarct [44]. In a study by Zhang et al. [21], adipose derived stem cells (ADSCs) transplanted into the myocar-dial scar tissue formed cardiac-like structures, induced angiogenesis and improved cardiac function. The APOLLO trial (NCT00442806) investigated transplanting fresh adipose derived MSCs to ST-elevated MI patients, and showed positive trends towards cardiac function, perfusion and neovasculogenesis (generally attributed to EPCs) [45]. The PRECISE trial (NCT00426868) looked at delivering adipose derived MSCs to patients with retractable angina, and noted no improvement in ejection fraction, but an increase in patient symptoms and exercise tolerance [46]. ANGEL is a Phase I trial that has completed enrolment for BioHearts Adipocell® therapy. Two Phase II studies have been initiated for adipose derived stem cells using intramyocardial injection; ATHENA (NCT01556022) for chronic myocardial ischaemia and MyStromal Cell (NCT01449032) [47] for chronic ischaemic heart disease and refractory angina where cells are pre-stimulated with vascular endothelial growth factor (VEGF). With regard to EPCs, early clinical studies have pointed to symptomatic benefits in patients with angina and cardiomyopathy [48-52]. In the ACT-34-CMI trial [49] investigators

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Fig. 2. Cell therapy: This figure shows different cell therapy approaches with different levels of sophistication and translational potential; unselected cells, purified cells and cells with materials.

assessed EPCs (or CD34 + cells) that were mobilised from bone marrow using granulocyte colony stimulating factor (G-CSF) for improving myocardial perfusion. The frequency of angina was significantly reduced compared to the control with the low-dose but not high-dose arms.

2.1.3. Skeletal myoblasts

Beginning almost 20 years ago, animal studies demonstrated that skeletal satellite cells or skeletal myoblasts showed promise in their ability to differentiate into myotubes or new myocardium and improve cardiac function post-infarction [53-60]. Skeletal myoblasts were transplanted from the skeletal muscle of a patient, purified, expanded and implanted into the heart [61]. The MAGIC trial revealed attenuation in LV remodelling, but no improvements in cardiac function, and was ultimately terminated due to increased risk of ventricular arrhythmias

[62]. The failure to improve myocardial function may be attributed to the inability of skeletal myoblasts to differentiate into cardiac myocytes

[63] or integrate electrically with the syncytium of the myocardium [63, 64]. Muscle derived stem cells [65] or cardiogenic muscle derived cell populations [66] may hold promise. MyoCELL® is a skeletal muscle myoblast cell therapy developed by B1OHEART [67] and is in Phase 11/ 111 trials in the US (MARVEL NCT00526253) in conjunction with the MyoCATH and MyoSTAR delivery catheters. Phase 1 trials and Phase 11 trials in Europe showed mixed results regarding increase in left ventricular ejection and clinical benefit [68-71 ].

2.1.4. Cardiac stem cells

Cardiac stem cells or CSCs are stem cells specific and resident to the heart. They are clonogenic, multipotent, self-renewing and can differentiate into three lineages; cardiomyocytes, endothelial cells and vascular smooth muscle cells. They express three cell-surface markers; MDR-1 (multi-drug resistant protein), C-kit (the receptor for stem cell factor), and/or Sca-1 (Stem cell antigen 1). Three methods for isolation of human cardiac stem cells have been described: (i) homogenizing large pieces of cardiac tissue and selecting CSCs using antibodies (usually limited to patients that undergo cardiac interventions such as bypass or transplant) [72], (ii) culturing a single biopsy and selecting CSCs with

antibodies as a subpopulation [73] and (iii) CSCs form cardiospheres 268 and can be selected by exploiting this property without the use of anti- 269 bodies [74]. CSCs reside in stem cell niches similar to those of highly 270 regenerating tissues in the post-natal senescent heart, and can undergo 271 symmetric or asymmetric division, giving rise to more CSCs or commit- 272 ted cells. When the heart tissue is injured, diseased or aged, resident 273 stem cell niches can also be affected, so the capacity of the heart to 274 self-heal is affected [75,76]. C-kit + progenitor cells are a candidate for 275 cell therapy and can be found in multiple species, and are reported to 276 be both essential and adequate for myocardial repair, without ruling 277 out participation of other cell types [77]. C-kit + cells have all the afore- 278 mentioned properties of cardiac stem cells, and were the first cardiac- 279 specific stem cell to be approved for a Phase 1 clinical trial SC1P1O 280 (NCT00474461) [78]. 1n the SC1P1O trial c-kit + cells were isolated 281 from a biopsy from the right atrial appendage taken during bypass sur- 282 gery and 1 million cells were delivered (mean of 115 days after M1) via 283 intracoronary injection to the infarction. 1nvestigators reported signifi- 284 cant increases in LVEFand decreases in a scar size of >30% [78,79]. How- 285 ever, this is an area of significant controversy in the literature, and 286 caution must be exercised with regard to the reported cardiogenic po- 287 tential of these cells. Recent work has reported that c-kit + cells can 288 only generate cardiomyocytes at a functionally insignificant level 289 (<0.03%), and that injection into diseased heart is unlikely to be respon- 290 sible for new cardiomyocytes [80]. Other work points towards the 291 concept that c-kit + precursors can generate cardiomyocytes in the 292 neonatal heart, but not the adult heart [81] or that in the neonatal 293 heart they are responsible for myocardial regeneration and 294 vasculogenesis, but in the adult heart they are only involved in 295 vasculogenesis [82], potentially explaining the reported clinical effects. 296 Another Phase 1 trial, CADUCEUS [83] examined the benefit of CSCs for 297 heart regeneration after myocardial infarction. C-kit + cells were har- 298 vested by an endomyocardial biopsy, and explants were cultured to 299 form cardiospheres [74,84]. Selected cardiospheres were infused into 300 the culprit arteries at 6 weeks to 3 months after M1 (1.25-2.5 x 107 301 cells). Scar size and left ventricular volumes benefitted from CSC thera- 302 py, but LVEF was not significantly increased. Follow-up studies have 303

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304 been initiated, and include RECONSTRUCT (NCT01496209) and ALLSTAR

305 (NCT01458405) for autologous and allogeneic CSCs, respectively.

306 2.1.5. Cardiopoietic stem cells

307 Directing the lineage of stem-cell populations towards specific or-

308 gans is promising, as cells can be obtained from more abundant sources

309 than the target organ itself. Additionally, risks associated with biopsy of

310 organs and issues with poor cell yields can be eliminated. Directing lin-

311 eage towards specific organs was originally described for pluripotent

312 embryonic stem cells [85-87], but can also be applied to adult stem

313 cell populations, including human MSCs. When exposed to certain

314 growth factors to upregulate cardiogenic potential, the cells are directed

315 down the cardiopoietic lineage [38,88]. The C-CURE trial investigates

316 delivery of cardiopoietic mesenchymal stem cells to ischaemic cardio-

317 myopathy patients. The trial demonstrated efficacy and safety of the

318 approach — with an increase in LVEF of 7% and positive effects on

319 haemodynamics and exercise tolerance [89]. Phase 111 trials CHART-1

320 and CHART-11 are starting in Europe and the US. These studies further

321 underline the trend towards pre-conditioning cells with growth factors

322 and even a hybrid approach where cells are delivered with growth fac-

323 tors or drugs, as discussed in the following section.

324 2.2. Additional considerations for cell therapy

325 Clinical translation needs to be the key consideration for cell therapy.

326 The optimal timing for cell administration and the effect of the extracel-

327 lular matrix must be fully understood. Studies are ongoing to elucidate

328 the mechanical changes in the infarct and mechanism by which the ex-Q6 tracellular environment of the infarcted area regulates the therapeutic

330 potential of stem cells. 1n a recent study researchers isolated and charac-

331 terized a diseased matrix to understand the effect of changes in infarct

332 stiffness over time on stem cell therapy [90]. Another factor for consid-

333 eration is the optimal endpoints for clinical trials. Many have used

334 ejection fraction as a metric of functional benefits, but whether this

335 translates into clinical benefits is not fully implicit and often doesn't cor-

336 relate with other functional parameters such as end systolic volume. A

337 metric of physical performance, such as the 6 minute walk test has

338 been included in recent trials, which makes sense, as the ultimate goal

339 of such regenerative therapy is to restore the patient's exercise toler-

340 ance and overall lifestyle to the pre-disease condition. Furthermore,

341 the timing of this type of functional testing is important, and in order

342 to evaluate the contribution of regeneration, a 6 minute walk test at

t1.1 Table 1

t1.2 Fold-increase in cell retention over intramyocardial saline delivery reported with various t1.3 injectable hydrogels.

t1.4 Study Hydrogel Time(s) Fold-increase in

of retention compared

analysis to saline control

4 weeks 1.3 for unaltered gel 15 pro-survival HGF included

24 h 1.3 (*NS)

5 weeks -2

48 h -2.5

4 weeks 2.5

90 min 1.77

24 h 1.5

4 weeks 8

24 h 1.75

4 weeks 2

1 day -1.5

1 week -1.9

2 weeks -2

4 weeks Presence of cells in chitosan group, none in control

t1.14 *HEC = Hydroxy-ethyl cellulose.

12 months should be employed to draw meaningful conclusions 343 (Fig. 2). Q7

2.3. Cells with biomaterial carriers 345

One of the major challenges in the clinical translation of cell therapy 346 is delivering and retaining viable cells in the heart tissue. The develop- 347 ment of cell therapy as a feasible therapeutic option is dependent on 348 methods to enable viable cells to reside in infarcted tissue and exert 349 therapeutic effects for extended periods. 1n cell therapy, isolated cell 350 suspensions in saline are usually administered systemically via intrave- 351 nous infusion or directly injected into the injured heart via the myocar- 352 dium, or perfused into the coronary arteries or veins. The cell therapy 353 clinical trials discussed in previous sections have primarily utilised 354 such simple cell delivery strategies. Saline solutions don't have the 355 capacity to localise and retain cells at the target site, and do little to 356 cater for the unique requirements of living cells with regard to provid- 357 ing biological cues to influence cell viability, behaviour and fate [8,33]. 358 Poor cell retention is likely to be a major contributing factor in the fail- 359 ure of cell-based therapies for M1 to achieve consistent and substantial 360 efficacy to date [3,91]. Among the possible mechanisms underlying 361 the phenomenon of poor retention are exposure of cells to ischaemia 362 and inflammation, mechanical washout of cells from the beating heart, 363 flushing by the coronary vessels, leakage of cells from the injection 364 site and anoikic cell death [92-94]. To address these issues there has 365 been a significant amount of preclinical research into material-based 366 cell therapy for cardiac repair. Delivered biomaterials can produce better 367 spatial distribution and potentially less problems with arrhythmogenicity 368 than simple saline injection techniques. A biomaterial scaffold can pro- 369 vide a surrogate ECM for encapsulated cells to enhance cellular viability 370 and enable physical retention at the infarct site. Biomaterials can pro- 371 vide protection from noxious insults like ischaemia and inflammation 372 and reduce cell death due to anoikis. Cell-loaded biomaterials address 373 the issue of mechanical dispersal of cells from the injection site, which 374 is a major source of cell loss within the myocardium and several studies 375 have shown that biomaterial delivery vehicles can enhance myocardial 376 cellular retention [95-97]. 1n short, biomaterials can help to deliver 377 more cells to the target site, keep cells localised and viable, and enhance 378 sustained production of beneficial paracrine factors at the target site. To 379 date, there exist two major biomaterial approaches to achieving cellular 380 delivery to the myocardium, namely cell-loaded injectable hydrogels 381 which encapsulate cells and polymerize in situ in the myocardial wall, 382 or preformed cell-seeded scaffolds which are affixable to the epicardial 383 surface [7], and both of these approaches will be addressed briefly here. 384

23.1. injectable hydrogels 385

Hydrogels can typically be injected via three routes: intracoronary, 386 epicardially or transendocardially. Such hydrogels have the potential 387 to rapidly exploit advancements in catheter technology for minimally 388 invasive delivery, reduced cost, shorter hospital times and potential for 389 multiple spatial and temporal administrations. To ensure injectability 390 the material and cells must facilitate loading into a catheter, the solution 391 must gel quickly at the site (but avoid premature gelation and catheter 392 blocking) and the gel must remain structurally sound for the course of 393 the therapy (to avoid embolization), and must degrade after cell thera- 394 py without producing toxic byproducts. The gel should also have 395 mechanical properties suitable for supporting the ventricular wall — it 396 must be robust, and endure the fatigue cycling of the heart throughout 397 the course of cell therapy. The increase in cell retention achievable can 398 become more dramatic over time. For example, Liu et al. reported a 399 1.5-fold increase in cell retention of adipose-derived stem cells encap- 400 sulated in chitosan/p-glycerophosphate/hydroxy-ethyl cellulose Q8 (chitosan//p-GP/HEC), 24 h post-administration via intramyocardial 402 injection, compared to cells delivered in saline [98]. However, an 8- 403 fold increase in retention was observed in hydrogel-injected animals 404 at day 28, which was likely related to a greater loss of cells from saline 405

tl.5 Zhang et al. [105]

tl.6 Yu et al. [106]

tl.7 Christman et al. [107]

tl.8 Habib et al. [108]

tl.9 Wangetal. [100]

tl.10 Martens et al. [109]

tl.ll Liuetal. [98]

tl.l2 Lu et al. [110]

tl.l3 Wang et al. [99]

PEGylated Fibrin + HGF

Alginate Microspheres Fibrin

PEG diacrylate PEG based Fibrin

Chitosan/ß-GP/*HEC Chitosan/ß-GP/*HEC Chitosan/ß-GP/*HEC

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Fig. 3. Cell-free therapy: Two types of cell-free therapy are discussed here; material based cell free-therapy and endogenous targeting, including RNA therapy, growth factors and proteins and small molecule therapy.

injected hearts over this time. A recent study shows that injectable chi-tosan not only improves retention of cells over time but also enhances cardiac differentiation of brown adipose derived stem cells and enhances functional improvements in the rat model [99]. Wang et al. used an a-cyclodextrin/poly(ethylene glycol)-b-polycaprolactone-(dodecanedioic acid)-polycaprolactone-poly(ethylene glycol) (MPEG-PCL-MPEG) hydrogel for bone marrow stem cell delivery, and showed improved retention in gel-injected animals, correlating with improved left ejection function and attenuation of scar expansion and left ventricular dilation, corroborating the hypothesis that biomaterial delivery can result in tangible enhancements in efficacy [100]. Collagen and laminin are the main components of myocardial extracellular matrix (ECM) and so can support cardiomyocyte attachment and elongation but the shape and dimensions of collagen and laminin biomaterial constructs have not yet been optimised. Future research may include designing 3-D shapes for these hydrogels, for example a collagen type 1 tubular scaffold has also been investigated [101], and shape memory injectable gels have been developed and should be considered for cardiac cell therapy [102,103]. An emerging technique for combining the advantages of hydrogel approaches with controllable, tailored tissue shape and size is bioprinting, enabling precise control over where cells are in the construct and the overall construct architecture to affect a particular cell fate or behaviour (Table 1) [104].

2.3.2. Preformed porous scaffolds

Porous or fibrous preformed scaffolds are the most common way for creating 3D constructs for cell delivery. In many cases, cells are grown on these constructs pre-implantation and patches are surgically attached to the epicardial surface. Leor et al. used a 3D alginate scaffold to construct a bioengineered cardiac graft in a rat model ofMI[111,112]

and subsequently optimised it for cell seeding and distribution. A colla- 435 gen patch was also used as a successful delivery vehicle for human mes- 436 enchymal stem cells and human embryonic stem cell derived- 437 mesenchymal cells for cardiac repair [113,114]. Cell attachment is an 438 important consideration in such constructs and they can be modified 439 with short peptides such as Arg-Gly-Asp (RGD); a peptide sequence de- 440 rived from the fibronectin signalling delay [115-119]. The selection, 441 density and patterning of binding sequences depend on the cell type 442 to be seeded on the matrix, and the natural ECM environment. Here, 443 we discuss porous scaffolds as carriers for cells to improve retention, 444 but a large volume of work has explored engineered heart tissue, so 445 the reader is referred to a comprehensive review [120] for more detail 446 on this. As an example, pre-conditioning of engineered heart patches 447 by cyclical mechanical stretch has shown to improve morphology and 448 contractile function of patches [121-128]. In a recent study electrospun 449 poly(e-caprolactone)/gelatin nanofibres were formed into a nanofibrous 450 patch to act as an improved method of cell retention (grafted MSCs 451 resulted in angiogenesis and facilitated cardiac repair) [129] as well as 452 providing mechanical support to the wall and acting as a ventricular 453 restraint, as discussed in the following section. The nanofibrous PG- 454 cell scaffold produced improvements in cardiac function (increase in 455 fractional shortening and ejection fraction, reduction in scar size and in- 456 crease in thickness in the infarcted area). Combinations of cell-loaded 457 gels and patches have been explored. Soler-Botija et al. describe prelim- 458 inary work on a fibrin loaded patch and an engineered bioimplant 459 (combination of elastic patch, cells and peptide hydrogel (Puramatrix, 460 Bedford, MA)) [130]. Electrical stimulation combined with 3D cell culti- 461 vation has also been explored. Nunes et al. describe the Biowire for plu- 462 ripotent stem cell-derived cardiomyocytes, consisting of a collagen gel 463 surrounding an electrically stimulated silk suture. These biowires had 464

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a stimulation rate-dependent increase in myofibril ultrastructural organization and conduction velocity (Fig. 3) [131].

3. Cell-free approaches

Cell-based strategies for cardiac repair involve delivering cells with potential for repair or regeneration to ischaemic or damaged areas of the heart. Despite the initial expectation regarding the cardiogenic potential of transplanted cells, in most studies the number of delivered cells that actually differentiate into cardiomyocytes is not large enough to account for observed clinical benefits, primarily due to low engraft-ment. The paracrine hypothesis may explain this, whereby released soluble factors from transplanted cells aid in regeneration [39,132]. There are a number of proposed mechanisms for such paracrine effects including increased angiogenesis, control of inflammatory responses, promotion of cardiac cell cycle re-entry and recruitment of endogenous stem cells, suggesting that paracrine targeting of endogenous cells may underlie many of the effects of cell therapy [41]. Similarly, delivery of cells has also been shown to produce mechanical reinforcement to the infarct scar area [133]. The field has undergone a paradigm shift, and investigators are renouncing the notion that therapy must be fixated solely around cells. Instead strategies such as acellular material-based approaches to produce mechanical reinforcement and tissue bulking in the myocardial scar and endogenous cell targeting through bioactive molecule delivery are subjects of extensive research to complement cell-therapy or to stand alone as cell-free therapy. Acellular strategies to cardiac repair have inherent advantages in that the lack of a required cell source could aid clinical translation.

3.1. Acellular material-based scaffolds

Material-based approaches target the important mechanical changes that occur post-myocardial infarction (or in chronic heart failure) resulting in ECM breakdown, geometric changes, LV dilation, stretched cardiomyocytes that can't contract, a growing borderzone and a spherical, thinning left ventricular wall [134-136]. Surgical ventricular restoration [137] (SVR), endoventricular circular patch plasty technique (Dor procedure) [138], partial ventriculectomy (Batista procedure) [139] and passive restraint devices such as the Acorn CorCap™ device [140,141], the Paracor Medical HeartNet restraint device [142], and the Myocor® coapsys device [143] all share the primary goal of reducing ventricular wall stress, and restoring left ventricular geometry. According to LaPlace's law T = P • R/t, where T, in this instance, is tension in the myocardial wall and varies proportionally to P (intraventricular pressure) and R (radius of curvature) and is inversely proportionally to t (myocardial wall thickness). By thickening the wall with a reinforcing material, stress can be decreased in the wall, especially around the infarct border zone [144]. Acellular injectable hydrogels and epicardial patches can be used to provide this tissue bulking wall reinforcement. If engineered to have specific biomechanical properties, this acellular material can promote the endogenous capacity of the infarcted myocardium to attenuate remodelling and improve heart function following myocardial infarction [145]. The elastic modulus can be tailored to match that of healthy myocardium or can be manufactured to have a higher elastic modulus to enhance tissue reinforcement [146], and numerical based simulations are valuable in predicting the response [144]. An optimal biomaterial should be able to balance the high forces that occur at the end of contraction in order to prevent or reverse maladaptive modelling [146]. The scaffold should be able to transfer the stress from the infarct-ed myocardium and border zone, and if the scaffold is biodegradable, cellular infiltration, vascularisation and formation of tissue should be sufficient to transfer the stress from the scaffold to the new myocardium before degradation. Injectable biomaterials used for acellular tissue reinforcement in animal models include fibrin [107,147], alginate [148-151], collagen [152], chitosan [98,110], hyaluronic acid [146, 153], matrigel [124,154], polyethylene glycol (PEG)-based materials

[155-157], acrylamides [158,159] and composites [160] of these materials. Both small animal studies [148,150] and large animal models [149, 160-162] have demonstrated benefit of this tissue bulking effect. For example, a biodegradable, thermoresponsive hydrogel for bulking of the ventricular wall based on copolymerization of N-isopropylacrylamide (NIPAAm), acrylic acid (AAc) and hydroxyethyl methacrylate-poly(trimethylene carbonate) (HEMAPTMC) was designed and characterized, and demonstrated an increase in wall thickness and capillary density, and ingrowth of contractile smooth muscle cells, thus offering a potential attractive biomaterial therapeutic strategy for ischaemic cardiomyopathy [158].

In addition to injectable materials, patches can be placed epicardially in order to provide wall thickening and reinforcement. Elastic patches such as polyester urethane urea have demonstrated an ability to produce an increase in fractional area change, and an attenuation of ventricular dilation in a rat MI model [163]. Engineered scaffolds or patches, such as a recently reported type 1 compressed collagen patch [145] can provide mechanical support to infarcted tissue, reducing dilation and fibrosis, increasing wall thickness and also increasing angio-genesis at the infarct zone and in the patch and border zone. This can lead to increased oxygen delivery and reduction in ischaemic tissue, and generation of new cardiomyocytes [145]. Clinically, an injectable hydrogel called Algisyl-LVR™ (LoneStar Heart, Inc., CA) has been used in a recently initiated Phase II trial AUGMENT-HF (NCT01311791). Circumferential intramyocardial injections of the alginate hydrogel remain in the heart (at the mid-ventricular level) as a permanent implant with the goal of increasing wall thickness, reducing wall stress and restoring ventricular geometry. Pre-clinical studies and a pilot study [164] show that the device has promise for decreasing ventricular volumes, increasing ejection fraction and wall thickness and decreasing myofibre stress at six months [164]. The AUGMENT-HF trial will evaluate the safety and efficacy of Algisyl-LVR™ as a method of left ventricular augmentation in patients with dilated cardiomyopathy, with a primary efficacy endpoint of change in peak VO2 (maximum oxygen uptake) from baseline to six months. This trial should provide some insight into the clinical benefits of the therapy. Another injectable alginate implant that has moved to clinical study is Bioabsorbable Cardiac Matrix (BCM), also known as IK-5001. After encouraging animal studies [148], recruitment is ongoing for PRESERVATION I (NCT01226563); a trial which investigates an in situ forming version of this hydrogel. An aqueous combination of sodium alginate and calcium gluconate is delivered in a bolus intracoronary injection, and into the heart muscle to form a flexible matrix that supports the heart physically and eventually dissipates and is excreted through the kidneys. The primary efficacy outcome measurement is left ventricular end diastolic volume index (LVEDVI).

The current limitations of acellular biomaterials are that optimal design parameters for therapeutic efficacy, including stiffness, degradation rate and bioactivities have yet to be determined. The experimental results in the literature reveal a complex biological and mechanical interaction between material and tissue. Experimental assessment of tissue bulking agents is mainly undertaken using a rat model of MI, which is not as clinically representative as a large animal model in terms of injection volume, injection method and volume of left ventricle. Injection time and data collection time also vary in these studies [165]. Further work is warranted to fully understand the specific mechanisms behind reported functional improvements. Only a small number of studies have directly compared different acellular biomaterials [166, 167], and the ideal acellular material properties have yet to be identified. It remains challenging to distinguish benefits resulting from changing the mechanical environment or benefits resulting from cardiac remodelling that is simultaneously occurring [168]. The in situ gelation rate of injectables must be rapid to avoid loss of material, but rapid gelation can make catheter delivery difficult. Lack of vascularisation in 3D scaffolds may also represent a limitation if scaffolds are intended for cell ingrowth and not just as a tissue bulking material. Cell survival may only be possible at the peripheries of 3D constructs, without

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593 vascularisation [169]. Furthermore, in the ischaemic human heart, there

594 may be a decreased production of factors that would promote vessel

595 sprouting. Provided tissue replacement is eventually envisaged, tissue

596 ingrowth and vascularisation must be sufficient for stress transfer to

597 newly generated myocardium before degradation, and the timing of

598 degradation to match tissue ingrowth will be critical to successful trans-

599 lation. 1f the purpose of the acellular biomaterial is to design an environ-

600 ment for endogenous cells to proliferate and regenerate, endogenous

601 cell numbers may not be high enough to initiate desired cell processes.

602 Acellular scaffolds cannot fully function as viable cardiac tissue replace-

603 ments, and are not fully biomimetic, potentially limiting the full poten-

604 tial of endogenous cells to recover through infiltration of the implant.

605 Acellular constructs negate the opportunity to pre-condition to enhance

606 functionality and integration with cardiac tissue. For example, cell-

607 loaded scaffolds can undergo mechanical and electrical pre-conditioning

608 that may result in a mature cardiac structure, higher force generation

609 and electrical coupling in the heart [122,127,128,170,171]. Although

610 true of all biomaterials, limitations for synthetic materials include difficul-

611 ties with scale-up of complicated chemical reactions and lack of innate

612 bioactivity, and with natural biomaterials limitations include difficulties

613 with regulatory approval and batch-to-batch variability [133]. Finally,

614 degradable materials can cause an inflammatory response and phagocy-

615 tosis [168], the effects of which are not fully characterized, and are cur-

616 rently reported to have beneficial [158] and counter-productive effects

617 [155].

618 3.2. Endogenous targeting

619 3.2.1. Small molecules

620 Small molecule drugs represent a promising therapeutic deliverable

621 for the treatment of ischaemic cardiomyopathy. These compounds are

622 often inexpensive to make and store. Advances in synthetic chemistry

623 mean that large libraries of structurally diverse molecules can be pro-

624 duced and screened for efficacy in modulation of a specific molecular

625 target. Similarly, a library of small molecules can be screened in a biolog-

626 ical system to determine novel drug targets and elucidate previously

627 unknown signalling systems implicated in myocardial disease. Struc-

628 ture activity relationship data can enable molecular modification to

629 optimise specificity, stability and efficacy. Such approaches are of dis-

630 tinct utility in clinical development. Small molecule drugs are currently

631 at an early stage of development for the purpose of myocardial regener-Q15 ation (for review see Jung and Williams [172]). Here, we discuss a con-

633 cise selection of candidate drug classes, with a particular focus on

634 advanced delivery to improve treatment outcomes.

635 3.2.1.1. Prostaglandins. Prostaglandins are endogenous small-molecule

636 fatty acid derivatives which mediate a variety of physiological effects.

637 Prostaglandin E2 and Prostaglandin 12 have a regenerative role in the

638 ischaemic myocardium and may have therapeutic potential post-M1.

639 3.2.1.2. Prostaglandin E2 (PGE2). Hsueh et al. demonstrated that daily

640 intraperitoneal administration of PGE2 enhanced cardiomyocyte re-

641 plenishment at the infarct border zone in a murine model of M1. Prosta-

642 glandin 12 (PG12) did not produce such effects, in this study. PGE2

643 increased the presence of Sca-1 + cells and regulated their potential

644 for a cardiomyogenic differentiation, suggesting that PGE2 could

645 activate and mobilise the endogenous CSC population. 1n addition,

646 PGE2 treatment rescued the ability of old mouse hearts to replenish

647 cardiomyocytes at the infarct border [173]. PGE2 is FDA approved for in-

648 duction of labour, and so possesses significant translational potential.

649 However, PGE2 is rapidly metabolised in vivo and so repeated dosing

650 was necessary in this study, which utilised a simple systemic route of

651 administration. This underpins the need for protective encapsulation

652 and delivery for long-term treatment and/or synthesis of more stable

653 prostaglandin mimics.

3.2.1.3. Prostaglandin 12 (PGI2). PG12 is a vasodilator and potent anti- 654

coagulant and has been FDA approved for the treatment of hyperten- 655

sion. Like PGE2, PG12 has a short half-life in vivo which is decreased in 656

conditions of myocardial infarction [174]. 1shimaru et al. delivered 657

ONO1301, a stable small molecule PG12 agonist on an epicardial colla- 658 gen patch to hamster hearts in a model of dilated cardiomyopathy 659

(but the observed therapeutic actions are likely also applicable to 660

acute M1), and found that ONO1301 treatment upregulated myocardial 661

expression of cardioprotective HGF, VEGF, SDF-1 and G-CSF. ONO1301 662 concentrations were found to be significantly higher in left ventricular 663

tissue than in systemic circulation for as long as two weeks after treat- 664

ment, highlighting the importance of local delivery and sustained 665

release. ONO1301 treatment preserved cardiac performance, in- 666

creased myocardial vascularisation, reduced fibrosis and prolonged 667

survival [174]. 1n a second study, Nakamura et al. encapsulated 668

ONO1301 in polylactic-co-glycolic acid (PLGA) microspheres which 669

produced a sustained release of drug for 10 days. The microspheres 670

were injected intramyocardially in a mouse model of acute M1 and 671

increased local HGF and VEGF expression, increased vascularisation 672

of the infarct border zone by day 7, decreased left-ventricular dilata- 673

tion and improved survival by day 28. ONO1301 was well tolerated 674

when delivered intramyocardially in PLGA microspheres. A Phase 1 675

clinical trial, where ONO1301 was administered orally was discontinued 676

due to diarrhoea in participants and systemic administration has been 677

shown to produce hypotension in experimental animals, highlighting 678

the importance of localised and controlled delivery in realising the full 679

potential of a given therapeutic strategy and avoiding off-target effects 680

[175]. 681

3.2.1.4. Pyrvinium Pamoate. Pyrvinium Pamoate (PP) is an FDA approved 682 anthelmintic drug, which inhibits NADH-fumarate reductase activity 683 essential for the anaerobic respiration of parasitic worms. Murakoshi 684 et al. postulated that the administration of PP could produce a differen- 685 tial cytotoxic effect in fibroblasts which proliferate in the myocardial 686 scar after infarct, and are reliant on anaerobic respiration in ischaemic 687 conditions, and hence enable anti-fibrotic therapy. PP was administered 688 orally, daily, beginning at 24 h after permanent left coronary artery liga- 689 tion (when the cardiomyocytes in the infarct area were likely dead) in a 690 mouse model of M1. There was a significant reduction in the presence of 691 fibroblasts in the infarct and border regions by seven days and fourteen 692 days and LVEF increased in PP treated animals. The authors also report 693 an increase in scar vascularisation, which they attribute to the permis- 694 sive microenvironment created by inhibition of fibrosis. PP therapy 695 was well tolerated [176]. 696

This is in contrast to a different study where Saraswati et al. admin- 697 istered PP via a single intramyocardial injection in a saline carrier at the 698 time of coronary artery ligation in a mouse model of M1, and observed a 699 significant increase in animal mortality upon PP treatment. 1t is likely 700 that administration of PP at this early stage enhanced cardiomyocyte 701 death in ischaemic conditions, resulting in larger infarcts and mortality, 702 and highlights the importance of time of dosage. Surviving animals did 703 not display a significant enhancement of cardiac regeneration or reduc- 704 tion of fibrosis. A once off injection of PP in saline may not have enabled 705 significant myocardial retention of the drug up to the time of initiation 706 of fibrosis. Therefore, utilisation of a biomaterial carrier, administered 707 at a minimum of 24 h post-infarct, which facilitated sustained release 708 may have ameliorated these results. Similarly, stimulus responsive 709 nanoparticles, tuned to deliver drug in a fibrotic environment or at the 710 time of initiation of fibrosis may have improved treatment outcome 711 [177]. While PP treatment was well tolerated when administered orally, 712 the risk for cytotoxicity to cardiomyocytes in the border zone where 713 perfusion is limited, or CSCs are naturally present in a hypoxic niche, Q16 may justify the use of targeted nanoparticulate carriers to ensure 715 increased specificity for fibroblasts and decreased risk for toxicity in 716 future studies [178]. 717

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3.2.1.5. Dipeptidylpeptidase IV (DPP-IV) inhibition. DPP-IV is a membrane bound peptidase which cleaves SDF-1. Pharmacological inhibition of DPP-IV aims to stabilise myocardial SDF-1 after MI, thereby enhancing recruitment of CXCR4 + circulating stem cells to effect regenerative efficacy. Zaruba et al. administered either Diprotin A, a small molecule DPP-IV inhibitor (twice daily systemic administration), G-CSF, to mobilise circulating progenitors or a multimodal administration of both in a mouse MI model. Combining G-CSF mobilisation and DPP-IV inhibition resulted in an increase in CXCR4 + cell homing to the myocardium, attenuation of infarct remodelling, neovascularisation in the infarct border zone, enhanced myocardial function and increased survival. Only the combination of Diprotin A and G-CSF treatment significantly attenuated myocardial remodelling, highlighting the potential of multimodal therapeutic strategies [179]. In addition, Theiss et al. demonstrated that a G-CSF/Diprotin A multimodal therapy significantly increased numbers of resident CSCs [180]. Given that it was necessary to administer Diprotin A twice daily to maintain efficacious concentrations within the myocardium, a sustained release formulation could greatly aid clinical translation. A Phase III clinical trial with another DPP-IV inhibitor, Sitagliptin, which has been approved for the treatment of hyper-glycaemia, in conjunction with G-CSF administration in patients with acute MI reported that the approach was well tolerated and appears feasible, but has yet to publish efficacy data [181].

3.2.2. RNA therapeutic strategies

3.2.2.1. Modified messenger RNA. A novel therapeutic strategy which has emerged recently is the delivery of modified messenger RNA (modRNA). Kormann et al. demonstrated that a collection of nucleotide modifications inhibited mRNA interaction with certain toll-like receptors, reduced immunogenicity and consequently enhanced stability when the modRNA was administered to mice. An intramuscular injection of modRNA produced a significant increase in target protein production in vivo. modRNA delivery to the lungs ameliorated a fatal genetic deficiency in mice despite only producing a very transient protein expression [182]. Warren et al. used modRNA delivery to create induced pluripotent stem cells, demonstrating that the transient expression of target proteins achievable could exert lasting effects on cell fate and differentiation [183].

Zangi et al. showed that modRNA encoding VEGF could transfect adult rat cardiomyocytes with a high efficiency (68%), using Lipofecta-mine, a commercially available transfection agent. The translational potential of Lipofectamine is unclear, however, since some authors have reported very low transfection efficiencies in large animal models or significant cytotoxicity in vitro [184,185]. One injection of modRNA/Lipo-fectamine transfected a significant portion of the mouse myocardium (25% of the left ventricle). Transgene expression peaked at 18 h and returned to baseline at 2-3 days, in contrast with DNA/Lipofectamine which peaked at 72 h and maintained high levels of expression for 10 days. VEGF modRNA/Lipofectamine was administered to infarcted mouse hearts, in comparison with VEGF plasmid DNA. Both VEGF DNA and VEGF modRNA increased vascular density in the infarct region but vessels produced by VEGF DNA were leaky, contributing to oedema which likely resulted in an observed increase in short-term mortality in VEGF DNA treated animals when compared to untreated controls. In contrast, modRNA VEGF treated animals showed decreased long-term mortality and improved cardiac function when compared to untreated controls, highlighting the importance of expression kinetics on functional outcome. VEGF modRNA treatment also upregulated Wt 1, an epicardial cardiac progenitor marker, in the infarct region, and in vitro data suggested that VEGF modRNA induced this cell type to undergo an endothelial differentiation, which may have contributed to treatment outcome [186].

The use of modRNA as a deliverable therapeutic confers several advantages over more conventional DNA therapy. Cytosolic expression avoids the risk of insertional mutagenesis associated with DNA therapy.

A transient, pulse-like expression more closely mimics endogenous paracrine signalling, in which sustained, high levels of expression over long periods, as produced with certain methods of DNA delivery, do not occur. Rather, a transient, strong signal, which is spatiotemporally controlled to act in the time and place it is required, is likely to be more efficacious and avoid undesired effects. While Zangi et al. has clearly demonstrated elements of this concept, further investigation into more clinically translatable nanoparticulate delivery vectors (as opposed to Lipofectamine) or localised therapy involving a biomaterial carrier will aid in unlocking the full potential of this technique. Such approaches may enable greater myocardial targeting and retention and spatiotemporal presentation of modRNA to maximise efficacy. modRNA therapy is currently in its infancy, and further investigation with other target genes to produce myocardial regeneration or offset the effects of ischaemic damage in vivo is warranted.

3.2.2.2. MicroRNA targeting. MicroRNAs (miRs) are endogenous, non-coding strands of RNA of around only 22 nucleotides in length. miRs are effectors of epigenetic regulation of protein expression, whereby a single miR demonstrates binding affinity for complementary oligonucleotide sequences in an array of mRNA targets, resulting in an inhibition of mRNA translation and/or mRNA degradation. Given that one miR typically has many mRNA targets, miR-mediated changes in protein synthesis are involved in a variety of complex intracellular signalling and modification of miR activity can have significant and multifaceted effects on cell phenotype.

miRs represent an attractive therapeutic target since they are extensively involved in cardiac development and postnatal disease processes including ventricular remodelling and fibrosis following infarction and processes with therapeutic applicability in acute infarction such as an-giogenesis or myocardial regeneration (for review see Fiedler and Thum [187]). Strategies to modify miR activity can take two forms — up-regulation of miR expression via transfection or viral transduction of target cells with a functional copy of a miR (a miR mimic), effectively inhibiting target protein expression, or inhibition of endogenous miR activity via complementary binding to synthetic anti-sense miRs or antagomirs, leading to an upregulation of target protein expression. Here, we highlight a concise selection of promising miR targeting strategies with different modes of action and discuss methods to enhance the delivery of miR to the infarcted heart.

Eulalio et al. undertook a high-throughput screening analysis of 875 miR mimics to identify 2 candidates (miR-590-3p and miR-199a-3p) which enabled the re-entry to cell cycle and proliferation of post-natal rat cardiomyocytes. These miRs were then delivered via intramyocardial injection of an adeno-associated viral vector to the infarcted mouse myocardium in vivo and significantly enhanced LVEF, increased wall thickness and reduced infarct size, primarily by stimulating cardio-myocyte proliferation [188]. Bonauer et al. demonstrated that miR-92a was expressed in endothelial cells and overexpression of this miR suppressed a variety of angiogenic processes in vitro. Conversely, a miR-92a antagomir enhanced angiogenesis in vitro and increased vascularisation of infarcted myocardium, reduced infarct size and enhanced cardiac function in a mouse model of acute MI, when administered intravenously. A panel of miR-92a target genes involved in vessel growth and development were identified [189]. Boon et al. determined that miR-34a demonstrated an increased expression in aged rat hearts which was related to age related decline in cardiac function. A miR-34a antagomir inhibited H2O2-mediated apoptosis in rat neonatal cardiomyocytes in vitro, and enhanced cardiac function, reduced cardiomyocyte apoptosis and enhanced vascularisation in a mouse model of acute myocardial infarction, when administered intramyocardially [190]. Hu et al. demonstrated that HL-1 cardiomyocytes transduced with miR-210 increased expression of pro-angiogenic growth factors and reduced caspase activity under hypoxic stress. When delivered intramyocardially via a minicircle non-viral vector in a mouse acute myo-cardial infarction model miR-210 reduced the presence of apoptotic cells

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and increased capillary density in the infarct area while enhancing left ventricular function. A panel of pro-angiogenic and anti-apoptotic miR-210 target genes were identified [191].

While these studies have demonstrated the preclinical potential of miRs for myocardial regeneration, significant hurdles to clinical translation remain. miRs not only represent a potentially powerful target to exert desired changes in cellular behaviour but also come with the risk of unpredictable off-target effects. Multiple target genes are controlled by a given miR, resulting in complex pharmacodynamics in both target and non-target tissues. miR delivery poses a challenge as unmodified miRs are rapidly degraded by systemic nucleases, may provoke an immune response and demonstrate low or unpredictable uptake by target cells. Significant modification of miRs to enhance stability has been achieved but sometimes at the cost of decreased specificity [192]. Therefore, targeted delivery of miR therapeutics to the myocardium utilising local delivery coupled with nanoparticulate and/ or biomaterial encapsulation is of the utmost importance.

The majority of studies investigating miR therapy for MI have used methods of miR delivery such as intramyocardial injection of viral vectors or simple systemic delivery of unencapsulated antagomirs. Such approaches not only provide a proof of concept for miR regenerative efficacy in the myocardium, but also pose translational hurdles such as safety concerns and lack of specificity for myocardial tissues. A range of nanoparticulate delivery vectors have been investigated for the targeted delivery of miRs, in a variety of different disease models outside of the cardiovascular field, with varying degrees of success and translational potential, including viral vectors, Poly(lactide-co-glycolide) (PLGA) particles, dendrimers, lipid based systems, Polyethylenimine (PEl)-based delivery systems and microvesicles such as exosomes (reviewed by Zhang et al., Muthiath et al. and Chistiakov et al. [193-195]). Gill et al. showed that ultrasound responsive microbubbles could transfect HL-1 cardiomyocytes with miR-133 upon application of ultrasound, which reversed cardiomyocyte hypertrophy. Such an approach could facilitate systemic delivery, but mediate miR uptake and expression only in tissues which are exposed to an externally applied ultrasound field [196]. Delivery of miRs in biomaterial carriers has also shown promise. Monaghan et al. determined that a collagen scaffold produced a sustained, bioactive release of miR-29B, which reduced maladaptive remodelling in a rat wound model [197]. In addition, local miR delivery in an injectable hydrogel has been shown to be an effective therapeutic strategy [198]. However, these approaches remain underexploit-ed in the field of miR therapy for myocardial regeneration and their future exploration may provide more translatable, safer and efficacious therapeutic strategies.

3.2.3. Direct reprogramming

A novel approach to effecting myocardial regeneration involves direct reprogramming of cardiac fibroblasts to functional cardiomyocytes or cardiac progenitor cells. Due to the limited regenerative potential of cardiomyocytes, the majority of the myocardial scar after Ml is composed of fibroblasts with no ability to contribute to the contractile activity of the myocardium. This technique involves therapeutic deliverables which aim to convert cardiac fibroblasts to cell types which can ultimately contribute to cardiac output. This has been investigated using several different approaches, including over-expression of cardiac transcription factors and delivery of microRNAs or small molecule drugs. Here, we discuss a concise selection of studies with a view to investigating clinical potential and suggesting scope for improvement using advanced delivery.

Recent research has identified sets of genes which, when over-expressed, can facilitate a direct reprogramming of cardiac fibroblasts to cardiomyocytes, while bypassing a pluripotent stem cell state (and the potential concomitant risk of tumour formation) [199]. Such transdifferentiation has been demonstrated in vitro [200] and has also shown clinical potential in vivo. Qian et al. reported that intramyocardial injection of three transcription factors, Mef2c, Tbx5, and myocardin

(GMT) encoded within retroviral vectors, resulted in minimal cardio-myocyte viral infection but significant transduction of fibroblasts in the myocardial border region of the infarcted mouse heart. 35% of cardiomyocytes in the infarct border zone were newly generated upon treatment and GMT delivery resulted in a decrease in infarct size and produced modest improvements in cardiac function [201]. Song et al. delivered GMT plus an additional factor, Hand2 (GHMT), via a retroviral vector through an intramyocardial injection in a mouse Ml model and determined that GHMT-treated animals had an LVEF of 49% compared to an untreated LVEF of 28%, which corresponded to twice the improvement of the controls and which persisted for up to 12 weeks [202].

Jayawardena et al. transfected murine cardiac fibroblasts with a combination of miRs 1,133,208 and 499 and reported transdifferentiation to a cardiomyocyte-like cell in vitro. The addition of a small molecule, JAK inhibitor 1, increased the efficiency of reprogramming 8-10 fold demonstrating the potential for small-molecule enhancement of this process. The miR cocktail was delivered intramyocardially via a lentiviral vector in a mouse model of Ml, and the results suggested that cardiac fibroblasts underwent a cardiomyocyte differentiation in situ but the authors did not investigate or report any potential effects treatment had on cardiac function [203]. In a recent study, Wang et al. utilised a small-molecule cocktail to reduce the number of genetic manipulations required to produce transdifferentiation of mouse fibroblasts to beating cardiomyocytes to just one — overexpression of Oct4. Cells passed through a cardiac progenitor stage during this transdifferentiation. Further development of this approach could lead to a fully pharmacological reprogramming, which could potentially circumvent some of the safety concerns of genetic manipulation. However, Wang et al. did not investigate this approach in vivo [199].

Clinical translation of fibroblast reprogramming techniques could be of significant therapeutic value. Direct reprogramming is a recent concept and consequently the majority of studies to date have served to provide a proof of concept, without significant focus on translational delivery approaches. As this field evolves, more clinically relevant delivery approaches and therapeutic deliverables will be explored. The use of viral vectors and stably expressed transgenes will likely pose transla-tional hurdles due to safety concerns. ln addition, the heart contains a large pool of fibroblasts, necessary for normal function [204]. It may be detrimental to target all cardiac fibroblasts non-selectively, and nanoparticulate targeting for fibroblasts present in or near the myocar-dial scar could aid in avoiding potential off-target effects of non-selective transdifferentiation. Such nano-particles could be responsive to stimuli in the scar environment itself, such as inflammation or reactive oxygen species, if a sufficient differential in molecular targets is not present between fibroblasts present in the scar and those elsewhere in the heart. Similarly, local delivery in biomaterial carriers could help to produce spatial control and retention of a therapeutic payload at the border zone.

3.2.4. Growth factors and proteins

Among the different therapeutic agents aimed to regenerate the damaged heart tissue after an ischaemic disease, peptides and proteins represent a well-consolidated acellular resource. The increased accessibility to these biopharmaceutical drugs and the advances in chemical modifications to enhance protein half-life in vivo and minimize immu-nogenicity [205] offer a broad range of new therapeutic modalities. Modified peptides and proteins can enable cardiac repair through activation of endogenous cardiac progenitor cells present at the injury site, the induction of cardiomyocyte proliferation and the recruitment of progenitor cells to damaged myocardium or of functional cells able to trigger neovascularisation.

With the aim to replace stem cell therapy in the treatment of acute myocardium ischaemic injury, Pavo et al. recently suggested the use of the secretome of apoptotic peripheral blood cells (APOSEC). The paracrine effects of this mixture of cytokines and growth factors were assessed after intramyocardial injection in a porcine model of acute

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MI. The administration of APOSEC produced downregulation of inflammatory and apoptotic genes 1 month after injection, whereas some angiogenic factors and regulators of vascular tone and homeostasis were upregulated. As a consequence, a reduced infarct size and improved hemodynamic function were found in APOSEC-treated animals [206].

Cell function is controlled by growth factors through the activation of specific signalling pathways [207]. The modulation mediated by these proteins may involve different biological routes and organs in the body. Therefore, the selection of cardiac-specific growth factors and safe dosing regimens should help prevent undesirable off-target effects. In the case of angiogenesis, vascular endothelial growth factor (VEGF) has been demonstrated to be a major regulator of vascularisa-tion under hypoxic conditions. As a potent growth factor for endothelial cells, VEGF administered after MI can induce angiogenesis and improve cardiac function. Despite its proven efficacy in preclinical models, VEGF has failed to achieve successful translation to clinical practice, in part due to dose limitation derived from the risk of nitric oxide-mediated hypotension [208]. Additionally, some concerns have been raised about the progression of metastatic tumour lesions as side effects of the prolonged administration of angiogenic growth factors.

The chemotactic stromal cell derived factor-l (SDF-I) has been described as a potent stem cell homing agent that is also involved in the regeneration of the vasculature. By binding to the CXCR4 receptor, SDF-I does not act as a growth factor on endothelial cells but increases the recruitment of endothelial progenitor cells [205]. This fact suggests a safer mechanism in the induction of angiogenesis since a therapy based on SDF-I may limit the uncontrolled formation of abnormal vessels. However, a major drawback of using SDF-I lies in its rapid cleavage by enzymes in the heart, such as DPP-IV and matrix metalloproteinases, leading to low efficacy. To surpass this disadvantage and improve its pharmacokinetics and activity, approaches based on altered SDF-I che-mokine designs that resist proteases or nanofibre-mediated delivery of SDF-I have been suggested [209]. In a complementary strategy, the conjugation of SDF-I to the soluble platelet collagen receptor glycopro-tein VI, which preferentially binds to collagen at exposed extracellular matrix in the damaged vasculature, enabled the targeted delivery of higher concentrations of SDF-I to the infarct site. This approach produced an enhanced recruitment of functional cells and a significant reduction of the infarct size in mice after MI [210]. Alternatively, gene transfer has been shown as a safe option in a Phase I clinical trial with a DNA plasmid encoding human SDF-I, JVS-100. The endomyocardial injection of the naked plasmid in patients with HF was well tolerated at all dose levels tested and led to improvements in clinical endpoints after 4 months [211].

Early clinical studies have also been performed with recombinant human neuregulin-l (NRG-I), a member of the epidermal growth factor family that promotes increased cell cycle activity and proliferation of cardiomyocytes through ErbB4 receptor binding. Patients with stable chronic HF showed an improved cardiac function with favourable acute and sustained hemodynamic effects after daily injections of NRG-I for eleven days [212]. Similarly to NRG-I, periostin can induce cell cycle reentry in adult cardiomyocytes. Kuhn et al. demonstrated that differentiated mononucleated cardiomyocytes have proliferative potential, and that periostin injected into the myocardium of rats after infarction has a regenerative effect, improving cardiac function after 12 weeks and reducing fibrosis and hypertrophy [213].

Hepatocyte growth factor (HGF) is a mesenchyme-derived pleiotro-pic factor with a stimulating effect on hepatocyte multiplication. Its implication in the regulation of cell growth, motility and morphogenesis of various cell types extends to the modulation of cardiovascular growth in pathological conditions. The antiapoptotic effect of HGF on cardio-myocytes has been demonstrated in rats after transient myocardial ischaemia and reperfusion [214]. Moreover, HGF may influence angio-genesis and progenitor cell recruitment. Urbanek et al. showed that a gradient of HGF facilitated translocation of CSCs from the atrioventricular

groove to the infarcted myocardium in mice [215]. A Phase II multicentre clinical trial evaluating a small-molecule mimetic of HGF, BB3, is currently ongoing with the aim to assess the safety of this drug in conjunction with standard care and its efficacy in improving heart function in patients following MI [216].

Growth and differentiation of recruited stem cells may be supported by insulin-like growth factor I (IGF-I). This hormone binds a tyrosine kinase receptor and enhances cell survival. IGF-I has been shown to reduce myocardial necrosis and apoptosis, and its overexpression in transgenic mice leads to an increase in myocyte turnover thus compensating for the extent of cell death in the ageing heart [217]. Moreover, in patients who had a diagnosis of ischaemic heart disease, low circulating IGF-I levels are associated with an increased risk in the development of cardiovascular disease [218]. The key role of IGF-I in cardiomyocyte homeostasis suggests a strong therapeutic potential. However, higher dose regimens have been associated with side effects such as hypotension and tachycardia. As proposed by O'Sullivan et al., a single local administration of low-dose IGF-I at 2 h into reperfusion may provide a prosurvival activity while avoiding significant side effects. In a porcine model of acute MI, the authors showed a reduced cardiomyocyte death at 24 h after IGF-I injection, which translated into structural and functional benefits in the regional and global myocardium 2 months after treatment [219].

In order to increase the bioavailability and control the release of growth factors in the cardiac tissue, drug delivery systems have been suggested as a means to protect and accumulate the protein cargo. Davis et al. reported the use of biotin-streptavidin to bind IGF-I to self-assembling peptides without interfering with bioactivity. These peptides provided a sustained IGF-I delivery for more than 1 month in rat myocardium. However, the co-injection of neonatal cardiomyocytes was necessary to achieve a therapeutic effect in rats after experimental MI [220]. To avoid the use of cell therapy, Chang et al. developed a delivery system based on PLGA nanoparticles functionalized with pPEI, which was able to electrostatically complex IGF-I. After comparing growth factor-loaded particles of different sizes (60 nm, 200 nm and 1 |jm), the authors found that the 60 nm-sized nanocarriers displayed the highest IGF-I activity in cultured cardiomyocytes. Following injection of these particles in the infarcted myocardium of mice, it was shown that the polymeric carriers prolonged IGF-I retention time and reduced cardiomyocyte apoptosis by more than 25%. Remarkably, a single administration of IGF-I-loaded nanoparticles improved cardiac systolic function, reduced infarct size and prevented ventricular remodelling at 3 weeks post-infarction [221].

The feasibility of controlled delivery using polymeric carriers was also shown for other proteins involved in repair of the damaged heart. Formiga et al. encapsulated FGF-l and NRG-I separately in PLGA micro-particles to assess the effect of cytokine sustained release on cardiac regeneration. The microparticle formulations showed very similar release kinetics with nearly 70% cumulative release within 1 month. The injection of the loaded particles into the ischaemic myocardium of rats produced reductions of the infarct size and fibrosis as well as an increase of the left ventricle thickness 3 months after treatment, with no significant differences among particles loaded with FGF-I, NRG-I or both [222]. In a different study with isolated rat cardiomyocytes in vitro, Johnson and Wang evaluated the protection from degradation and the sustained release of the morphogen Sonic hedgehog (Shh) from a coacervate delivery system [223]. Shh is known to control the epithelial/ mesenchymal interactions during the embryonic development, and has demonstrated potential to restore blood flow in a mouse model of hindlimb ischaemia after multiple injections for 1 month [224]. The formulation of Shh-heparin complexes in poly(ethylene argininylaspartate diglyceride) prolonged the release of Shh for over 3 weeks and provoked an upregulated secretion of VEGF, IGF-I, SDF-I and Shh by cardiac fibroblasts for at least 2 days.

As an alternative to particle formulations, the encapsulation of proteins in carrier gels also provides a controlled release and enhances

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12 C.L. Hastings etal./ Advanced Dru;

retention in the target area. 1n a rabbit model of M1, Fujita et al. showed efficient angiogenesis and collateral flow induced by FGF-2 loaded in photocrosslinkable chitosan hydrogels. The chitosan aqueous solution containing FGF-2 was applied on the surface of the ischaemic myocardium and subsequently crosslinked by UV-irradiation for 30 s. Notably, the chitosan hydrogel allowed an extended delivery of FGF-2 for a period longer than 1 month [225]. An ideal growth factor carrier should have the ability to flow through a catheter, enabling minimally invasive application, and thereafter form a solid gel to avoid the injected drugs to be pumped out of the heart. 1n an attempt to develop such system, Wu et al. synthesized a biodegradable aliphatic polyester hydrogel, poly(6-valerolactone)-poly(ethylene glycol) (PEG)-poly(S-valerolactone), which gels when heated at physiological temperature. The injection of the hydrogel in the infarcted myocardium of rats attenuated adverse cardiac remodelling and improved ventricular function for up to 35 days. These effects were strengthened by covalently attached VEGF, which additionally provided increased regional angiogenesis in comparison with free VEGF co-injected with the hydrogel [226]. With the same aim to design an injectable biomaterial, Bastings et al. proposed pH-sensitive ureido-pyrimidinone PEG hydrogels, which are fluid above pH 8.5 and instantaneously gel at neutral pH. By transcath-eter injection of the synthetic hydrogel incorporating both HGF and 1GF-1 in a porcine model of M1, the authors demonstrated a safe administration and a reduction in scar collagen after 1 month [227].

Tissue regeneration is often characterized by complex cascades of growth factors with critical roles in cell proliferation and differentiation. The combination of several growth factors is required to mimic the native environment and promote the formation of functional tissue [208]. Since myocardial repair involves the contribution of different signalling pathways, the combined activation by co-administered growth factors represents a promising approach for an enhanced performance of CSCs and may also enable effective and safe angiogenic interventions.

Ellison et al. demonstrated the superiority of co-administered HGF and 1GF-1 to induce myogenic differentiation of endogenous porcine CSCs in the presence of adult rat ventricular myocytes in vitro. The injection of a small dose of 1GF-1 and HGF through the coronary artery supplying the infarcted region in pigs produced a dose-dependent protective effect on myocardial survival and reduced hypertrophy in the peri-infarct zone. Furthermore, a reduced infarct size and enhanced left ventricular function were measurable 2 months after the treatment [228]. 1n a different approach, Song et al. recently reported the combination of SDF-1 with the angiogenic tetrapeptide Ac-SDKP to activate regenerative mechanisms in a model of chronic HF in rats. The authors immobilized Ac-SDKP in acrylated hyaluronic acid hydrogels, in which SDF-1 was added before crosslinking. 1nterestingly, hydrogels with single SDF-1 or Ac-SDKP failed to show a significant regenerative activity whereas the dual therapy led to increased angiogenesis, improved left ventricular function, decreased infarct size and higher wall thickness at 4 weeks after hydrogel injection [229]. 1n spite of these promising preliminary results, more extensive knowledge on the role of different stem cell homing factors and the potential synergies with differentiation and proliferation mechanisms is needed. As exemplified by some negative reports on the use of SDF-1 therapies for M1 in vivo [230], a tight control of the complex molecular signalling is likely required to avoid unexpected effects.

Temporal control on the release of proteins is another key factor to realise their maximal potential for cardiac regeneration. 1n the case of granulocyte colony stimulating factor (G-CSF), which induces proliferation of haematopoietic stem cells with the capacity to regenerate the in-farcted myocardium, an effect was found only in patients who received G-CSF early after M1 [231 ]. As hypothesized by Ruvinov et al., a sequential delivery of 1GF-1 and HGF may favour the regenerative process: a fast release of 1GF-1 could enhance survival of the remaining functional myocardium, while a more sustained release of HGF could induce angiogenesis and more favourable remodelling at later stages. By bioconjugating 1GF-1 and HGF individually with alginate-sulphate, and combining both

ivery Reviews xxx (2014) xxx-xxx

complexes with low viscosity sodium alginate solution, dual-release 1175 injectable hydrogels were obtained. The intramyocardial injection of 1176 the alginate gels in a rat model of acute M1 produced an increased 1177 cytoprotection and angiogenesis in the infarct after 1 month when 1178 compared to the administration of 1GF-1 and HGF in saline. Furthermore, 1179 the sequential treatment induced a higher level of cell proliferation at 1180 the infarct border after 1 week, as well as a higher expression of 1181 GATA-4 after 4 weeks, indicative of angiogenesis, survival and stem 1182 cell recruitment [232]. 1n another example, albumin-alginate microcap- 1183 sules were employed to separately incorporate FGF-2 and HGF with 1184 different release kinetics. As the authors of this study suggest, the se- 1185 quential release of FGF-2, which generates a potent angiogenic activity, 1186 followed by the arteriogenic signalling induced by HGF, resulted in a 1187 mature vessel network that prevented cardiac hypertrophy and fibrosis 1188 and led to improved cardiac perfusion after 3 months in a rat model of 1189 chronic HF [233]. 1190

Furthermore, a time-controlled combination of immune response 1191 inhibition and neovascularisation was recently achieved by Projahn 1192 et al. By crosslinking thiol-functionalized copolymers of ethylene 1193 oxide and propylene oxide with different agents, i.e. hydrogen peroxide 1194 or PEG-diacrylate, the authors obtained degradable gels with disulphide 1195 or thioether bonds, respectively. 1n the presence of reduced glutathione, 1196 the disulphide-based gels degraded in 1 day (fast degradable hydrogel, 1197 FDH) while complete degradation of thioethers occurred after 1 month 1198 (slow degradable hydrogel, SDH). On the one hand, an inhibitor of neu- 1199 trophil infiltration, MetCCL5, was released from FDH to block the 1200 immune response during the first hours. On the other hand, SDF-1 was 1201 released from SDH for a sustained recruitment of haematopoietic stem 1202 cells. The co-administration of both loaded hydrogels in the infarcted 1203 myocardium of mice preserved cardiac function, promoted angiogene- 1204 sis and facilitated wound healing processes [234]. 1205

Together with fibroblast growth factors, bone morphogenetic proteins 1206 (BMPs) and wingless-type (Wnt) proteins are involved in the initial 1207 specification of cardiac cells. Yoon et al. showed that the combination 1208 of BMP-2 with FGF-4 induced myogenic differentiation of MSCs 1209 in vitro, and that the implantation of MSCs treated with the growth fac- 1210 tors enhanced engraftment and myogenic differentiation in infarcted 1211 myocardium in rats [235]. BMP-2 has been demonstrated to improve 1212 the contractility of individual spontaneously beating cardiomyocytes. 1213 Moreover, intravenous injection of BMP-2 in a mouse model of acute 1214 M1 induced a reduction in cardiomyocyte apoptosis up to 4-fold in the 1215 border zone and up to 2-fold in the remote myocardium when com- 1216 pared to negative controls 5 to 7 days after administration [236]. 1n 1217 the case of Wnt, Duan et al. found that Wnt1 and Wnt7a were signifi- 1218 cantly upregulated after acute cardiac injury. The expression of Wnt1 1219 peaked within 2 days after injury and was sustained at lower levels 1220 for two weeks, driving an early repair response in mice myocardial 1221 ischaemia [237]. 1t has been demonstrated that the Wnt1/p-catenin sig- 1222 nalling system mediates a pro-fibrotic repair in cardiac fibroblasts after 1223 M1. A close correlation of the responsiveness of cardiac fibroblast to Wnt 1224 and the temporal pattern of Wnt1 expression after heart injury suggests 1225 the role of this pathway during cardiac disease [238]. 1226

1n addition to its main role in haematopoiesis, erythropoietin (EPO) 1227 presents antiapoptotic and pro-angiogenic properties that have shown 1228 efficacy against M1 in different animal models. 1n rats, intraperitoneal 1229 administration of EPO once every 3 weeks induced new vessel forma- 1230 tion associated with enhanced mobilisation, myocardial homing and 1231 vascular incorporation of endothelial progenitor cells. Accordingly, 1232 VEGF levels increased 4.5-fold in the groups treated with EPO [239]. 1233 Kawachi et al. showed that subcutaneous injection of EPO enhanced an- 1234 giogenesis in pigs following M1 by upregulating HGF and FGF systemi- 1235 cally and VEGF and 1GF in the border and infarct areas [240]. Despite 1236 substantial evidence of EPO effectiveness in vivo, clinical studies failed 1237 to show expected therapeutic efficacy [241 ]. As suggested by Roubille 1238 et al., meta-analysis of the available data from clinical trials could help 1239 in assessing the impact of factors such as the route of administration 1240

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Towards Advanced Delivery

Fig. 4. The case for advanced delivery, as discussed here, is summarised by four main concepts; localised therapy, nanoparticle encapsulation, minimally invasive delivery and multimodal approaches.

1241 or the timing of EPO treatment [242]. To facilitate the clinical translation

1242 of the cardioprotective role of EPO found in animals, larger clinical trials

1243 with consistent inclusion/exclusion criteria might be needed.

1244 Given increasing knowledge on the different molecular pathways in

1245 which growth factors and cytokines are involved and new develop-

1246 ments in biopharmaceutical drug combinations to maximise therapeu-

1247 tic potential, enhanced treatment options for cardiac regeneration are

1248 expected to occur in the coming years. In addition, the formulation of

1249 these therapeutic agents in drug delivery systems will facilitate a safer

1250 administration and more effective dosing patterns, leading to improved

1251 clinical outcomes.

1252 4. The case for advanced delivery

1253 Regenerative therapy for ischaemic cardiomyopathy is an extremely

1254 active area of research and a variety of potential treatment strategies

1255 have emerged over recent decades. Cell therapy has arguably

1256 progressed furthest towards clinical translation, as evidenced by a

1257 significant number of clinical trials, but is still hampered by poor and

1258 unpredictable efficacy when implemented in large patient cohorts.

1259 Indeed, translation of the positive results achievable in preclinical

1260 models has been largely slow and unsatisfactory for all avenues of myo-

1261 cardial regenerative therapy. With this in mind, we elected to review a

1262 selection of therapeutic approaches with a particular focus on advanced

1263 delivery strategies as a method to enhance efficacy, reduce deleterious

1264 effects and aid clinical translation. These concepts are summarised here.

1265 1. Localised therapy in biomaterials — this encompasses the local deliv-

1266 ery of therapeutic agents in biomaterial carrier vehicles as opposed to

1267 simple systemic delivery. This is of particular importance for cellular

1268 payloads where a biomaterial can act to mimic the natural ECM, to

1269 enhance survival and provide biological cues for cellular behaviour

1270 and fate. In addition, the localised delivery of small molecules or

1271 growth factors within a biomaterial matrix permits for sustained re-

1272 lease over extended periods to enhance efficacy in target tissues.

1273 2. Nanoparticulate encapsulation — this involves the delivery of thera-

1274 peutics in a nanoparticulate carrier to reduce interaction with off-

1275 target tissues and enhance targeting to the ischaemic myocardium.

1276 3. Multimodal approaches — the concurrent delivery of more than one

1277 therapeutic (for example cells with small molecule drugs) can

1278 achieve synergistic efficacy. Release of therapeutics from either an

1279 implantable biomaterial or nanoparticle system can also be tailored

1280 to mimic a biological cascade. For example, sequential release of

1281 two or more agents can be utilised to target early and late stage effi-

1282 cacy in a physiological process such as angiogenesis [243].

4. Minimally invasive delivery approaches — percutaneous catheter 1283

systems can be utilised to locally deliver therapeutic agents to the 1284

heart in a minimally invasive manner, reducing surgical time and 1285

cost, and allowing multiple administrations of therapy. 1286

The first two concepts have been addressed in the context of the pre- 1287 vious sections and the following section will focus on the latter points, 1288 discussing the potential of these delivery approaches in the pursuit of 1289 clinical translation and improved treatment outcomes. In particular, 1290 we will discuss the potential for multimodal therapeutics primarily in- 1291 volving the combination of cells with an additional co-delivered thera- 1292 peutic, and the state of the art with regard to minimally invasive 1293 catheter delivery to the myocardium. 1294

4.1. Multimodal therapeutic strategies 1295

A multimodal combination of cells with an additional therapeutic 1296 agent represents a particularly attractive therapeutic strategy. This 1297 approach confers the potential for therapeutic agents to act on co- 1298 delivered cells, as well as exert efficacy in target tissues. Co-delivery in 1299 a biomaterial carrier can ensure that both cells and a second therapeutic 1300 deliverable are kept in close proximity for the duration of therapy to 1301 enhance synergistic interaction (Fig. 4). Q22

A number of studies have addressed the potential of co-delivering 1303 cells with growth factors to produce therapeutic angiogenesis, which 1304 could be of significant utility in the treatment of ischaemic cardiomyop- 1305 athy. The hindlimb ischaemia model is often used to gauge the potential 1306 of a given therapeutic strategy to produce vascular growth. For example, 1307 Saif et al. administered PLGA microparticles containing a triple combi- 1308 nation ofVEGF, HGF and Angiopoietin-1 (Ang-1) alone, human cord 1309 blood vasculogenic progenitor cells (ECFCs) alone, or a combination of 1310 both, via intramuscular injection in a murine hindlimb ischaemia 1311 model. Cells or growth factor loaded particles alone produced a modest 1312 increase in vascularisation and limb perfusion but a multimodal combi- 1313 nation produced a substantial further increase. The biomimetic ratio- 1314 nale was to combine two potent pro-angiogenic agents, VEGF and 1315 HGF, with a vessel pro-maturation agent, Ang-1. This was proposed to 1316 avoid the phenomenon of leaky and poorly functional vessels which 1317 can in some cases occur upon treatment with VEGF alone. In an ear tis- 1318 sue leakage assay, the authors showed that administration of VEGF 1319 alone produced significantly leaky vessels, which was somewhat ame- 1320 liorated by co-administration of HGF and significantly reduced by triple 1321 administration ofVEGF, HGF and Ang-1. The triple combination also 1322 produced more vessels than VEGF/HGF co-administration, highlighting 1323 the importance of multimodal administration and biomimetic strategies 1324 to enhance efficacy [244]. 1325

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t2.1 Table 2

t2.2 Comparison of commercially available cell injection catheters by access, core needle outer

t2.3 diameter, material and shape.

t2.4 Device Manufacturer/research group Needle shape

t2.5 Endocardial delivery

t2.6 Helix BioCardia Helical

t2.7 MyoCath Bioheart Straight, can be deflected

t2.8 MyoCath II Bioheart Weeping

t2.9 C-Cath® Cardio3 Biosciences Curved, large-to-small side holes

t2.10 Myostar Bioheart Straight

t2.11 Stiletto Straight

t2.12 t2.13 Transvascular

t2.14 TransAccess Medtronic Curved

t2.15 Cricket/Bull-Frog Mercator Medical Straight, mounted on balloon

t2.16 t2.17 Epicardial

t2.18 Cell-Fix Chachques group Straight, attached to "sucker" fixation system

t2.19 t2.20 Intracoronary perfusion

t2.21 PTCA devices Multiple No needle, cells delivered

through guidewire lumen

Multimodal combinations of cells and growth factors have also been investigated in the infarcted myocardium. Dvir et al. investigated the delivery of neonatal rat cardiac cells on an alginate patch containing bound IGF-1, SDF-1 and VEGF to act as a co-delivered pro-survival and pro-angiogenic cocktail. The patch was prevascularised on the omen-tum before implantation on the infarcted rat heart. Patches containing growth factors demonstrated enhanced vascularisation on the omen-tum, and prevascularised patches produced greater myocardial regeneration in terms of increase in left ventricular function and reduction in ventricular remodelling, although patches containing no growth factors were not investigated in the infarcted heart [245].

Padin-Iruegas etal. injected self-assembling peptide nanofibres with tethered IGF-1 (NF-IGF-1) alone, rat CPCs (rCPCs) or a combination of both in a rat myocardial infarct model, with the rationale that co-delivered IGF-1 would increase delivered cell survival along with enhancing the regenerative response of resident CPCs. Both CPCs and NF-IGF-1 were injected intramyocardially and NF-IGF-1 facilitated presentation of bioactive IGF-1 for a sustained period. Combination therapy produced greater enhancement in LVEF, increased the presence of newly formed cardiomyocytes (230% compared to NF-IGF-1 alone),

and increased infarct vascularisation and reduction in infarct size, with Q23

respect to the delivery of cells or IGF-1 nanofibres alone. In addition, 1347

combination therapy enhanced the activation of resident CPCs [246]. 1348

Takehara et al. administered bFGF in a gelatin hydrogel 1349

sheet alone, human cardiosphere derived cells (hCDCs) alone, or a 1350

multimodal combination of both to the infarcted porcine myocardium 1351

via intramyocardial injection (hCDCs) or surgical implantation on the 1352

epicardium (hydrogel sheet). Sustained release of bFGF from the gel- 1353

atin sheet for up to three weeks was achieved. Delivery of bFGF/gel- 1354

atin alone enhanced myocardial perfusion and LVEF while hCDCs 1355

alone enhanced LVEF and reduced infarct volume. Co-delivery of 1356

hCDCs and bFGF/gelatin significantly enhanced hCDC engraftment in 1357

the myocardium and resulted in synergistic increases in LVEF and 1358

reductions in infarct size, compared with delivery of either hCDCs 1359

or bGF/gelatin alone. No synergistic effects were observed when bone- 1360

marrow-derived hMSCs were co-delivered with bFGF, supporting the 1361

hypothesis that cardiac-derived stem cells are likely more suited for 1362

cardioregenerative applications [247]. 1363

On the basis of these promising results this approach (CSC/bFGF 1364

therapy) has progressed to a small Phase I clinical trial, ALCADIA (AutoL- 1365

ogous human CArdiac-Derived stem cell to treat Ischemic cArdiomyop- 1366

athy) to determine the safety of the approach. Autologous CSCs were 1367

administered to patients via intramyocardial injection and bFGF/gelatin 1368

sheets were implanted epicardially, during bypass surgery. Patients 1369

demonstrated increased LVEF and reduced infarct size after the surgical 1370

procedure, but in the absence of a control group and as a result of a small 1371

patient cohort, definitive conclusions about efficacy were not possible. 1372

The trial demonstrated that the approach was safe and feasible and fur- 1373

ther trials will establish the efficacious potential of this approach [248]. 1374

In an interesting acellular hybrid therapy approach Kubota et al. 1375

employed an atelocollagen sheet/polyglycolic acid ventricular restraint 1376

device (VRD) alone, a small molecule PGI2 agonist ONO1301 on an 1377

atelocollagen sheet alone, or a multimodal ONO1301-doped VRD in a 1378

canine model of myocardial infarction. At 8-weeks post-infarction 1379

hearts treated with the multimodal VRD demonstrated the greatest 1380

increase in LVEF, greatest reduction in left ventricular wall stress and 1381

ventricular remodelling. All hearts treated with ONO1301 (either 1382

alone or in combination with VRD) demonstrated an increase in myo- 1383

cardial vascularisation and upregulation of HGF, VEGF and SDF-1 in 1384

the myocardium [249]. In a similar hybrid approach with cells, Shafy 1385

et al. showed that the combination of adipose-derived stem cells 1386

(injected into the infarct and seeded in a collagen matrix) with a poly- 1387

ester CorCap VRD device resulted in significant improvements in Q24

Transvascular Delivery

Fig. 5. Current access routes for cell-based therapies to the heart include transvascular delivery, intracoronary perfusion, epicardial delivery and endocardial delivery. An example of a device designed for each delivery route is depicted in this figure.

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ejection fraction, systolic function and diastolic function in a sheep infarct model [250]. This semi-degradable ventricular bioprosthesis approach is an example of biomaterial-mediated cell therapy combined with a constraint device. The CELLWAVE study addressed delivery of BM-MSCs combined with a pretreatment of low energy cardiac shockwave to improve honing of cells and expression of SDF-1 and VEGF. The combination of shock wave with cells resulted in an increase in ejection fraction of 3.2% [251]. Chachques has bioengineered nano-biomaterials with elastomeric membranes to acquire a controlled drug release patch to which they can tailor for local cell attraction and cell differentiation [252].

Multimodal approaches show particular promise for myocardial regeneration. However, the biomedical industry is sometimes reluctant to pursue such therapeutic strategies due to the concern that it could result in a longer regulatory process and consequent delays in bringing a product to market. Multimodal therapeutics can be more difficult to classify and categorise since they involve a variety of therapeutic elements. However, the enhanced potential for improved treatment outcomes and therefore a product with a greater chance of obtaining clinical approval means that multimodal approaches should receive serious consideration for future therapies. This is especially true given the lack of concrete clinical translation in this field to date, despite decades of research, primarily into simplistic treatment approaches involving systemic delivery of single agents or cells. The FDA opened an Office of Combination Products in 2002, specifically to provide guidance to clarify the regulation of combination therapies and to enable timely and effective premarket review of combination products [253]. In addition, preclinical and clinical safety and efficacy data for pre-existing single agent regenerative therapeutics are likely relevant to new combination product applications, reducing the overall regulatory burden.

4.2. Minimally invasive therapy — catheter delivery

It is important that deliverable therapeutic formulations reach the region of the infarcted myocardium where they are most required. The heart resides in the thoracic cavity and in general is accessed via highly invasive surgical procedures involving a thoracotomy, contributing to significant costs and patient morbidity. In order to facilitate localised delivery to the myocardium in a minimally invasive way, percutaneous catheter delivery can be employed. Percutaneous catheters are medical devices which generally consist of flexible, hollow tubing and an associated guide wire with a distal 'active' tip which performs an injection. The device can be passed into the vasculature through a small incision, advanced and manipulated via a proximal handle, until the tip reaches the therapeutic target.

Catheter delivery of cells alone, typically in a saline carrier, has been more explored than catheter delivery of more advanced materials such as patches or hydrogels, and will be discussed briefly here. The trans-catheter cardiac cell delivery field has recently been directed at improving cell retention. In contrast to thoracic surgical injections or patch implantations, transcatheter approaches are less invasive. They allow the effect of cell therapy to be evaluated independently of other surgical procedures, and justify multiple deliveries of cells. The following sections will describe existing delivery systems, their capabilities, and will suggest potential for innovation in areas where suitable devices are not commercially available. For a more detailed insight into current systems the reader is referred to two review papers on this area [254, 255]. Several catheter-based access approaches have been used in humans; directly injecting cells into the ventricular wall (epicardial, endocardial and transvascular approaches), and infusing cells into the coronary arteries using existing balloon angioplasty catheters [254, 255]. Table 2 and Fig. 5 describe a panel of available devices. The delivery systems differ in their access approach, but share some common features; a low profile core element dedicated to transport cells, which has a bevelled needle to anchor into the myocardium, and outer components to protect the core and deliver it to the infarcted tissue.

The endocardial delivery devices approach the myocardium from inside the ventricle. As for many interventional cardiology catheteriza-tions, they are introduced to the arterial system transfemorally or transradially, guided around the aorta, and through the aortic valve in a retrograde fashion. Catheters are manipulated inside the ventricle by support catheters or steerable designs, and can rely heavily on imaging systems for accurately targeting injection sites at ischaemic areas or the infarct border zone. Transvascular devices approach the myocardium from the epicardial surface. A support catheter is placed through the femoral veins, and tracked around to one of the coronary veins. By using an IVUS (IntraVascular UltraSound) system, the nearby coronary artery and the pericardium can be localised. The coronary vein is then punctured with a small needle, and the injection catheter is passed through this puncture site to the epicardial wall. For epicardial access, the Cell-Fix catheter includes a retractable needle and a polyurethane umbrella shaped suction system which fixes the device to the epicardi-um when connected to vacuum. This allows stability for penetration and retraction of the injection needle [256]. The goal of intracoronary infusion is to increase the number of cells delivered to the ischaemic myocardium. Vessels are visible by angiography techniques and if cells are injected proximally, they can be distributed to large areas of the myocardium. The method uses established interventional cardiology tools such as Percutaneous Transluminal Coronary Angioplasty (PTCA) devices, where the cells are delivered through the guidewire lumen on removal of the guidewire when the device has been steered through the vasculature to the culprit vessel. Limitations include the fact that large cells in viscous suspensions may not be appropriate due to the risk of obstruction, and cells used must be capable of migrating across the endothelium to perivascular spaces. Furthermore, if patients have chronic total occlusion, this approach is not feasible. PTCA catheters are not designed or approved for cell infusion, and there are no standard tests to compare them for this purpose. Early studies with these devices reported low retention of cells from direct injection, retrograde venous delivery and intracoronary perfusion groups, albeit with slightly higher numbers for the direct injection group [68,257]. More recently, analytical and numerical modelling based on the Darcy Law and transport mass retention has led to optimised needle designs specifically for cell retention [258]. The use of a small-to-large graded side-hole design in a 75° curved Nitinol needle in the C-Cath lessened interstitial pressure during delivery to improve retention and resulted in a significant (>3-fold) increase in cell retention (healthy and infarcted hearts) [258]. While the catheters described here are a huge improvement on simple systemic or invasive local delivery, they are limited in that they are only optimised to deliver a simple saline payload which doesn't facilitate sustained release or cell viability; there is still a need for catheters delivering retentive materials such as injectable hydrogels or epicardial patches.

4.2.1. Catheters for material based approaches

Existing catheter technology may not be appropriate for injecting hydrogels due to considerations such as rapid gelation kinetics, hydro-gel viscosity and complications with gelation triggers such as thermal sensitivity or requirements for mixing and incorporation of crosslinking agents immediately prior to injection. Additionally, there is a lack of available devices for catheter-based delivery of preformed scaffolds, patches or cell sheets. For injectable hydrogels, certain catheter design criteria need to be fulfilled to maintain the liquid prepolymer during catheter transit to the injection site, to allow fast gelation in situ once the polymer has been injected, and to provide multiple deliveries without issues such as needle blockage. New cyto-compatible catheterized devices such as double-barrel injectors (to mix chemically crosslinked gel precursors with crosslinking agents), cooled catheters (for thermo-responsive gel payloads) and epicardial patch deployment tools are needed. Several preclinical studies have determined the feasibility of delivering injectable hydrogels to the heart using commercially available catheter systems. For example, Leor et al. delivered an alginate

1516 1517Q27

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hydrogel to the coronary vessels in pigs using an injection catheter [149]. Martens et al. determined the optimum viscosity and gelation parameters for a fibrin hydrogel for use with a range of commercially available catheters [109]. Singelyn and Christman determined that an in-situ gelling decellularised myocardial matrix was compatible with catheter delivery [91]. Other groups have improved conventional catheters or syringes for their purposes; Kofidis et al. describe a Y-shaped applicator for two syringes where the matrix is contained in one syringe and cell suspension in the other whereby homogeneous mixing occurs on injection [154].

Until now, delivery of patches or scaffolds in preclinical trials has been performed in a surgically invasive manner during open chest procedures. Patches are still largely delivered to the epicardium, due to concerns of embolization upon endocardial deployment. The field of epicardial delivery could learn lessons from other interventional fields, such as that of Total Aortic Valve Replacements (TAVIs) and other procedures using the transapical delivery approach. This access route could be a promising candidate for epicardial material mediated-delivery. In this approach, access to the epicardium is undertaken via a mini-thoracotomy and a pericardial incision. Device profile is only limited by the constraints of the pericardial space, therefore design constraints of transapical access catheters are not as limiting as transvascular catheters when delivering a material that requires a higher profile catheter bore. These tangible design targets and the significant amount of research in the evolving field of material based therapy are compelling reasons for innovation in minimally invasive delivery systems for material-based cardiac regenerative therapy. Finally, ventricular restraint devices can be combined with cells, biomaterials or endogenous targeting approaches. Clinical trials have investigated the delivery of cells while patients are receiving left ventricular assist devices (for example the ASSURANCE trial NCT00869024), and the hybrid approach of ventricular unloading with cell delivery has shown promise for improving native cardiac function, allowing removal of mechanical assistance and potentially obviating the need for a heart transplant [259-262]. Future promising work will focus on combining cells with extra-cardiac assist devices for biomaterial-based cell delivery on assist device implantation with multiple follow-ups, consisting of minimally invasive cell administrations (a cell 'top-up' dose) via transvascular catheter delivery. Local delivery of biomaterials via catheter systems could reduce the time, invasiveness and cost of a given therapeutic procedure while capitalising on the pro-retentive, cytocompatible and sustained release properties of biomaterial therapeutic formulations. Future development of such systems might greatly aid clinical translation of cardiac regenerative strategies.

4.3. Conclusion

Advanced delivery strategies are of the utmost importance in fully realising regenerative therapies for the treatment of ischemic cardiomy-opathy. Simple delivery of cells, growth factors or drugs has shown promise, especially pre-clinically. However, clinical translation remains elusive. Physiological and pathological processes in the heart are inherently complex, and consequently more sophisticated therapeutic strategies which fully utilise advanced delivery techniques may be required to enable clinical translation. The preclinical evidence presented in this review suggests that an ideal therapeutic might utilise a combination of the discussed delivery approaches. This strategy might involve minimally invasive catheter delivery of a biomaterial carrier vehicle. The implanted biomaterial bolus should ideally contain a multimodal pay-load consisting of cardiac-derived stem or progenitor cells combined with biomaterial-encapsulated nanoparticles. Such nanoparticles should facilitate a controlled release of bioactive molecules which exert therapeutic efficacy on co-encapsulated cells and local tissue for sustained periods. Alternatively, bioactive molecules could be free-loaded into the biomaterial matrix, provided a sustained release is possible. The formulation should seek to maximise myocardial retention and uptake. Where possible, the formulation should seek to emulate endogenous

biological cues and processes to maximise efficacy, through judicious alteration of design criteria such as duration and sequence of bioactive molecule release, spatial presentation of implanted therapeutics and manipulation of encapsulated cell behaviour and fate. If systemic delivery is required, it should be undertaken using targeted nanoparticles to enhance drug accumulation in myocardial tissue and reduce off target effects.

Acknowledgements

We acknowledge the funding derived from the European Union's Seventh Framework Programme for research, technological development and demonstration under grant agreement no NMP3-SME-2013-604531 from November 2013 to October 2017, entitled AMCARE, and the Fulbright International Science and Technology Fund.

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